Bioceramics Volume 10 (Biocermaics)

Bioceramic s Volume 10 Edited by Lauren t Sedel Laboratoire de Recherches Orthopediques Faculte de Medecine Lariboisie...

Author: L. Sedel | C. Rey

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Bioceramic s Volume 10

Edited by Lauren t Sedel Laboratoire de Recherches Orthopediques Faculte de Medecine Lariboisiere-Saint Louis Universite Paris 7 - Denis Diderot Paris, France

Christia n Rey Ecole Nationale Superieure de Chimie Institut National Polytechnique de Toulouse Toulouse, France

PERGAMON

Pergamon/Elsevie r Titles of Relate d Interes t

BIOCERAMICS BOOK SERffiS Bioceramic s 4: W. BONFELD, G. W. HASTINGS and K. E. TANNER (ISBN 0 7506 0269 4) Bioceramic s 6: P. DUCHEYNE and D. CHRISTIANSEN (ISBN 0 08 042143) Bioceramic s 7: O.H. ANDERSSON, R.-P. HAPPONEN and A. YLI-URPO (ISBN 0 08 042144 X) Bioceramic s 8: J. WILSON, L. L. HENCH and D. GREENSPAN (ISBN 0 08 0426778) Bioceramic s 9: T. KOKUBO, T. NAKAMURA and F. MIYAJI (ISBN 0 08 0426840) RELATED JOURNAL Biomaterials(ISSN 0142-9612) For details on the above Pergamon/Elsevier Science book series or a free specimen copy of any Elsevier Science journal please contact your nearest Elsevier Science office (see copyright page for addresses).

Bioceramic s Volume 10 Proceedings of the 10th International Symposium on Ceramics in Medicine, Paris, France, 5-9 October, 1997 Editedby Lauren t Sedel Christia n Rey

U.K.

Elsevier Science Ltd, The Boulevard, Langford Lane, Kidlington Oxford 0X5 1GB, U.K.

U.S.A.

Elsevier Science Inc., 655 Avenue of the Americas, New York 10100, U.S.A.

JAPAN

Elsevier Science Japan, Tsunashima Building Annex, 3-20-12 Yushima, Bunkyo-ku, Tokyo 113, Japan

Copyright ' 1997 Elsevier Science Ltd All Rights Reserved No part of thispublication may be reproduced, stored in a retrievalsystem or transmittedin anyform or by any means: electronic, electrostatic, magnetic tape, mechanical, photocopying, recording or otherwise, without permission in writing from the publishers. First Edition 1997 Librar y of Congres s Catalogin g in Publicatio n Dat a A catalog record for this book is available from the Library of Congress Britis h Librar y Cataloguin g in Publicatio n Dat a A catalog record for this book is available from the British Library ISBN 0 08 0426921

Cover picture BOURDALOU Factory of Comte d’ Artois Limoges - End of XVIII century Adrien Dubouche National Museum, Limoges

Printed in Great Britain at the University Press, Cambridge

Organizin g Committe e Chairme n

Laurent Sedel Christian Rey

Member s

Alain Meunier Didier Bernache-Assollant

Secretar y

Martine Henry-Amar

Scientific Committe e D. Bernache-Assollant J. Delecrin P. Frayssinet M. Hamadouche P. Marie A. Meunier A. Moroni

M. Nardin R. Nizard N. Passutti H. Petite S. Redey L’H. Yahia

Internationa l Advisor y Committe e A. Barbosa (Portugal) P. Boch (France) W. Bonfield (UK) G. Daculsi (France) P. Ducheyne (USA) P. Griss (Germany) K. de Groot (Netherlands) U. Gross (Germany) L.L. Hench (USA)

S.F. Hulbert (USA) T. Kokubo (Japan) H. Oonishi (Japan) J.A. Planell (Spain) A. Ravaglioli (Italy) C. Rey (France) J. Wilson (USA) T. Yamamuro (Japan) A. Yli-Urpo (Finland)

Scientific Endorsemen t European Society for Biomaterials Society for Biomaterials (USA) European Society for Orthopaedic Research Japanese Society for Biomaterials Ceramics Society of Japan Japanese Society of Orthopaedic Ceramic Implants National Institute for Scientific and Medical Research (I.N.S.E.R.M., France) National Centre for Scientific Research (C.N.R.S., France)

This Page Intentionally Left Blank

Prefac e The 10th Bioceramics volume reports on the meeting held in Paris from 5th to 8th October 1997. It has been a great pleasure for me to edit this book with the highly respected scientist, Christian REY. Bioceramics in Medicine has become one of the major fields in biomaterials. Many different bioceramics, with various material compositions and characteristics producing different biological behaviour, result in a wide range of medical applications. Dense ceramics, such as aluminium oxide, present a range of properties which should allow the material to last for many years within the living body. This highly oxidised material should also theoretically resist corrosion related to oxygen capture. The latter exhibits very interesting tribological properties when compared to either metal or polyethylene, which are more widely employed. Bioceramics made of calcium phosphate of different chemical compositions texture and porosity, are of great interest for the replacement of defective bone lost through tumour excision or major trauma or to enhance bone repair after fracture, fusion or ridge augmentation. Calcium phosphate ceramics have been employed to copy bone mineral. The imitation of bone mineral is very difficult as it presents a complex system within which formation, degradation and even texture are not fully understood. However, artificial calcium/phosphate of various compositions and even textures has been proved to be bone friendly. Osteoconduction has been demonstrated by many of these materials and their capacity to provide a scaffold for bone cells is also well recognised. Is this related to the three dimensional aspects of these materials and to the ability of mesenchymal stem cells to grow on this support? Or is osteoconduction only related to the chemical composition of the ceramic? Even if the answers to these questions are not available at the present, results confirm the excellent osteoconductivity of calcium/phosphate ceramics. Further questions need to be addressed: what are the events following bone apposition on the surface of the ceramic? Will this material resorb allowing natural bone to replace the artificial material, providing either an excellent bone union or bone apposition. Or will this material remain unchanged providing a permanent bone apposition or bone augmentation necessary for a shelf procedure ? Bioglasses, containing different types of silica with phosphorus, represent another very interesting field. These materials bond to bone without any visible interface. Do these bioglasses provide enough bone bonding strength? Do they enhance bone union? Experimental results appear to be positive, but confirmation in the human

viii Preface

body is yet to come. Mixed materials such as apatite wollastonite glass ceramics (AWGC) show both remarkable tolerance as well as strength and an ability to remain unchanged for many years, when implanted in the living body. As well as clinical applications, fundamental subjects are also discussed. What is the mechanism of bone cell osteoconduction? Is it related to chemical or physical factors, or both? What is the role of the micro and the macro geometry of the surface? Would it be possible to imitate these characteristics on other materials, such as metals or polymers? All these questions and many others are discussed in this book. The applications of bioceramics are numerous. Orthopaedic and dental surgery are the two major fields of interest but others such as plastic surgery, E.N.T., problems of percutaneous devices, embolisation materials, calcium phosphate cement are also increasingly important. Finally the bioceramics field is one of the most active biomaterials applications and has grown considerably in the last twenty years. Its high level scientific input from a wide range of backgrounds and the participation of industries create a great interest in this subject. We hope that everybody, whatever his own interest, will find something of value in this book.

Laurent SEDEL

CONTENTS CALCIU M PHOSPHAT E IN VIVO FORMATIO N Precipitatio n of Calcium Phosphat e on Titani a Ceramic s K-L. Eckert, S.-W. Ha, S. Ritter and E. Wintermantel Apatit e Formatio n on Polymer s by Biomimeti c Processin g Using Sodium Silicate Solution F. Miyaji, S. Handa, T. Kokubo and T. Nakamura Bonelike Apatit e Formatio n on the Surfac e of ChemicaU y Treate d Tantalu m Substrates : Effect of Heat Treatmen t T. Miyazaki, H.M. Kim, F. Miyaji, T. Kokubo and T. Nakamura

3

7

11

Influenc e of Hydroxyapatit e Particl e Size and Morpholog y on Hapex^ ^ M. Wang, R. Joseph and W. Bonfield

15

Effect of Fluorid e Substitutio n on the Biocompatibilit y of Hydroxyapatit e K.A. Ring, L. Di-Silvio, I.R. Gibson, C. Ohtsuki, L.J. Jha, S.M. Best and W. Bonfield

19

Apatit e Precipitatio n in Biphasi c Calcium Phosphat e Cerami c After Incubatio n in Rabbi t Serum and Ionic Simulate d Body Fluid (SBF) R. Rohanizadeh, J.M. Bouler, D. Couchourel, M. Padrines and G. Daculsi

23

Apatit e Precipitatio n in Biphasi c Calcium Phosphat e Cerami c After Implantation : Influenc e of Implantatio n Site R. Rohanizadeh, M. Trecant-Viana, J. Delecrin, J.M. Bouler and G. Daculsi

27

GLAS S CERAMIC S BIO ACTIVIT Y Bioactivit y and Structur e of Organicall y Modified Silicate Synthesize d by the Sol-Gel Metho d K. Tsuru, S. Hayakawa, C. Ohtsuki and A. Osaka Hydroxyapatit e Formatio n on Bioactive Glass Coate d Titaniu m C.Y. Kim and S. Kwon

33

37

X Contents

Effect of Multivalen t Cation s in Calcium Silicate Glasses on Bioactivit y N. Imayoshi, C. Ohtsuki, S. Hayakawa and A. Osaka

41

Transformatio n of Bioactive Glass Granule s into Ca-P Shells In Vitr o S. Radin, P. Ducheyne, S. Falaize and A. Hammond

45

Multilayere d Coating s of Hydroxyapatite/Glas s Cerami c Composite s Plasm a Spraye d on Ti-6A1-4V Alloy P.L. Silva, J.D. Santos and FJ. Monteiro

49

The Bony Reactio n to Rapidl y Degradabl e Glass-Ceramic s Based on the New Phas e Ca2KNa(P04) 2 C. Miiller-Mai, G. Berger, C. Voigt, B. Bakki and U. Gross

53

Resorbable , Porou s Phosphat e Inver t Glasses - Firs t in Vitr o and in Vivo Result s J. Vogel, K.-J. Schulze, D. Reif, P. Hartmann, U. Platzbecker and B. Leuner

57

Implantatio n of Bioactive and Iner t Glass Fibre s in Rat s - Soft Tissue Respons e and Short-Ter m Reaction s of the Glass M. Brink, P. Laine, K. Narva and A. Yli-Urpo

61

Spina l Fusion Using Titaniu m Spacer s With Bioglass^ and Autogenou s Bone: A Comparativ e Stud y in Sheep J. Wilson, G. Lowery and S. Courtney

65

DENSE AND POROU S BIOACTIV E CERAMIC S Macroporou s Biphasi c Calcium Phosphat e Ceramics : Influenc e of Macropor e Diamete r and Macroporosit y Percentag e on Bone Ingrowt h O. Gauthier, J-M. Bouler, E. Aguado, P. Pilet and G. Daculsi Mechanica l Fatigu e of Hot Presse d Hydroxyapatit e S. Raynaud, E. Champion, D. Bernache-Assollant Mechanica l Propert y Change s in Macroporou s Cerami c After Implantatio n into Bone and Muscle M. Trecant-Viana, J. Delecrin, J.M. Nguyen, J. Royer and G. Daculsi

71

75

79

Contentsxi

Differenc e of Bonding Behavior Between Four Differen t Kinds of Hydroxyapatit e Plat e and Bone S.S. Chung, C.K. Lee, K.S. Hong and H.J. Yoon

83

Treatmen t of Osteomyeliti s by Antibiotic-Soake d Porou s A-W Glass Cerami c Block K. Kawanabe, Y. Okada, H. lida and T. Nakamura

87

Calcium Hydroxyapatit e Cerami c Implant s Impregnate d With Antibioti c for the Treatmen t of Chroni c Osteomyeliti s Y. Yamashita, T. Yamakawa, K. Kato, Y. Shinto, N. Araki and A. Uchida

91

BONE CELL S ONT O BIOACTIV E CERAMIC S Measuremen t of Intac t Osteocalci n Content s in the Composit e of Porou s Hydroxyapatit e Cerami c and Allogeneic Marro w Cells M. Akahane, H. Ohgushi, T. Yoshikawa, S. Tamai, Y. Dohi, K. Hosoda and T. Ohta Si-Ca-P Xerogels and Bone Morphogeneti c Protei n Act Synergisticall y on Rat Stroma l Marro w Cell Differentiatio n In Vitr o E.M. Santos, P. Ducheyne, S. Radin, B. Shenker and L Shapiro Effect of Surfac e Instabilit y of Calcium Phosphat e Ceramic s on Growt h and Adhesion of Osteoblast-Lik e Cells Derived from Neonata l Rat Calvari a T. Suzuki, M. Hukkanen, L.D.K. Buttery, J.M. Polak, Y. Yokogawa, K. Nishizawa, F. Nagata, Y Kawamoto, T. Kameyama and M. Toriyama Histologica l Evaluatio n of Culture d Bone Graf t Using Cryopreserve d Marro w CeUs H. Nakajima, T. Yoshikawa, H. Ohgushi, M. Akahane, S. Tamai, K. Mishima and K. Ichijima Interaction s of Bioceramic s on Huma n Osteoarthriti s (OA) Type B Synoviocytes. Effects on Interleuki n Levels and Lipoxygenas e Pathway s B. Liagre, J-L. Charissoux, M-J. Leboutet, D. Bernache-Assollant and J-L. Beneytout A Long Term Implantatio n of Culture d Bone in Porou s Hydroxyapatit e T. Yoshikawa, H. Ohgushi, H. Nakajima, M. Akahane, S. Tamai and K. Ichijima

97

101

105

109

113

117

xii

Contents

SURFAC E BEARIN G CERAMIC S Oxide Ceramic s for Articulatin g Component s of Tota l Hip Replacement s G. Willmann

123

Ex Vivo and In Vitr o Analysis of the Alumina/Alumin a Bearin g System for Hip Join t Prosthese s H.J. Refior, W. Plitz and A. Walter

127

Hybri d Alumina-Alumin a Hip Replacement : A Survivorshi p Analysis and Result s at a Minima l Five Year FoUow-Up M. Hamadouche, P. Bizot, R.S. Nizard and L. Sedel

131

Low Temperatur e Ageing Behaviour of Zirconi a Hip Join t Head s J. Chevalier, J.M. Drouin and B. Cales Characterizatio n of Zirconi a Coate d by Bioactive Glass: Preliminar y Observation s M. Bosetti, M. Santin, M. Mazzocchi, A. Krajewski, M. Rastellino, A. Ravaglioli and M. Cannas

135

139

Calcium Phosphat e Particle s are Found at the Polyethylen e Inser t Surfac e Whethe r Implante d With Ha-Coate d Devices or Not. A SEM-EPM A Study P. Frayssinet, L. Gineste, G. Bonel and N. Rouquet

143

Bone Remodellin g Aroun d Implante d Material s Under Load-Bearin g Condition s M. Oka, Y.S. Chang, S. Yura, K. Ushio, J. Toguchida and T. Nakamura

147

CLINICA L USE OF CERAMIC S Clinica l Comparativ e Stud y Between Porou s Coate d and Hydroxyapatit e Porou s Femora l Implant s Y.H. Kim, J.H. Shon and I.Y. Choi

153

Revision Rate s and Radiographi c Change s Associated With Differen t Socket Interfac e Technologie s : Clinica l Result s from 416 Patient s at 6 To 8 Year s FoUow-Up M.T. Manley, A. Edidin, J. A. Epinette, R.G. Geesink, J.A. D’Antonio and W.N: Capello

157

Contents xiii

Comparativ e Stud y of the Result s Between Custom Non-Coate d Cementles s Hip Implant s and Mirrore d Cementles s HA-Coate d Hip Implant s on the Contra-Latera l Side M. Mulier and G. Deloge

161

Improvemen t of THR With Spongiosa Meta l Surfac e Using the Wear Couple Ceramic-On-Cerami c G. Quack, G. Willmann, H.G. Pieper and H. Krahl

165

Acetabula r Reconstructio n in Revision Tota l Hip Arthroplast y Using a Bone Graf t Substitut e R.P. Pitto and D. Hohmann

169

POSTE R 1 Effect of Solution Ageing on Sol-Gel Hydroxyapatit e Coating s B. Ben-Nissan, C. Chai and K.A. Gross Ionic Cement s : Influenc e of the Liquid/Soli d Rati o on Porosit y and Mechanica l Propertie s F. Betchem, P. Michaud, F. Rodriguez and Z. Hatim

175

179

Sinterin g and Therma l Decompositio n of Hydroxyapatit e Bioceramic s J. Cihlar and M. Trunec

183

An Elaboratio n of the New Dissolution Mechanis m for Apatit e S.V. Dorozhkin

187

Ultrastructura l Study of Long Term Implante d Ca-P Particulat e into Rabbi t Bones A. Dupraz, R. Rohanizadeh, J. Delecrin, P. Pilet, N. Passuti and G. Daculsi

191

Bioactive Glass-an d Glass-Cerami c Composite s and Coating s M. Ferraris, E. Yerne, A. Ravaglioli, A., Krajewski, L. Paracchini, J. Vogel, G. Carl, C. Jana

195

Mechanica l Characterisatio n of Bioactive Coating s on Zirconi a E. Verne, M. Ferraris, C. Moisescu, A. Ravaglioli and A. Krajewski

199

xiv

Contents

Vacuu m Plasm a Spraye d Titaniu m and Hydroxyapatit e Coating s on Carbo n Fiber Reinforce d Polyetheretherketon e (Peek )

203

S.-W. Ha, A. Gisep, H. Gruner, J. Mayer and E. Wintermantel Low Temperatur e Crystallizatio n of Hydroxyapatit e Sputtere d Films in an Autoclav e

207

J. Hamagami, K. Nakamura, Y. Sekine, K. Yamashita and T. Umegaki Fabricatio n of In-Cera m Cor e by Sheet Formin g Proces s

211

D-J . Kim , M-H . Lee and C.E . Kim Surfac e Structur e of Bioactive Titaniu m Prepare d by Chemica l Treatmen t

215

H.M. Kim, F. Miyaji, T. Kokubo, T. Suzuki, , F. Itoh, S. Nishiguchi and T. Nakamura Prefabricate d Biological Apatit e Formatio n on a Bioactive Glass-Cerami c Promote s In Vitr o Differentiatio n of Feta l Rat Chondrocyte s

219

C. Loty, S. Loty, T. Kokubo, N. Forest and J.M. Sautier The Effect on Mechanica l Propertie s by Osteoblasti c Cell Ingrowt h in Macroporou s Syntheti c Hydroxyapatit e and Interpor e 200 ^ ^

223

E. Nordstrom, H. Ohgushi, H. Yoshinari, S. Tamai and T. Yokobori Characterizatio n and Cell Reactio n of a-TCP - and HAp-Coating s on Titaniu m Plat e

229

M. Ohgaki, S. Nakamura and M. Akao The Ectopi c Osteoconductio n Model

233

H. Ohgushi, M. Okumura, T. Yoshikawa, H. Ishida, H. Yajima and S. Tamai Condition s of the Coprecipitatio n of Calciu m Hydroxyapatit e With Zr02 , ZrO i " Y2O3, AI2O3 from Aquoeu s Solution s Using Ammoni a

237

V.P. Orlovskii, Zh A. Ezhova and E.M. Koval Transformatio n of a-TC P to Hydroxyapatit e in Organi c Media

241

K. Sakamoto, S. Yamaguchi, A. Nakahira, K. Kijima and M. Okazaki Structur e and Solvation Effects of P04^- , HP04^- , H2P04 ’ and H3PO 4 from AMI and PM 3 A.J. Sahnas, A. Serret, M. Vallet-Regi and L.L. Hench

245

Contents xv

The Detaile d Configuratio n of Carbonat e Ions in Apatit e Structur e Determine d by Polarize d Ir Spectroscop y

249

Y. Suetsugu, I. Shimoya and J. Tanaka Tissue Cultur e on Amorphou s Calciu m Phosphat e Coatin g

253

K. Suzuki, Y. Kageyama, Y. Kita, A. Yoshino, K. Matsushita and T. Kokubo Bonelike Apatit e Layer Forme d on Organi c Polymer s by Biomimeti c Proces s : TEM-ED X Observatio n of Initia l Stag e of Apatit e Formatio n H. Takadama, F. Miyaji, T. Kokubo and T. Nakamura Sinterabilit y and Second Phas e Formatio n of Syntheti c Hydrox y Apatit e : Controllin g Parameter s and Effect on Bond Strengt h

257

261

H-J. Youn, H.S. Ryu, K.S. Hong, S.S. Chung and C.K. Lee Porou s Sol-Gel Bioglassfi from Near-Equilibriu m Dryin g

265

J. Zhong and D.C. Greenspan

POSTE R 2 Ceramic-Cerami c Bearin g System s Compare d on Differen t Testin g Configuration s

271

J. ChevaHer, B. Cales, J.M. Drouin and Y. Stefani Dissolution and Mechanica l Behaviou r of Plasma-Spraye d Cerami c Coating s for Orthopaedi c Application s

275

N. Demonet, P. Benaben, J.L. Aurelle, B. Forest and J. Rieu Design of a Calciu m Phosphat e Bone Cemen t Suitabl e for the Fixatio n of Meta l Endoprosthese s

279

FCM. Driessens, L Khairoun, MG. Boltong and Ja. Planell Quantitativ e Compariso n of In Vivo Bone Generatio n With Particulat e Bioglassfi and Hydroxyapatit e as a Bone Graf t Substitut e

283

Y. Fujishiro, H. Gonishi and L.L. Hench Test of Bioactivit y in Four Differen t Glasse s A.M. Gatti, L.L. Hench, E. Monari, F. Gonella and F. Caccavale

287

xvi

Contents

Compariso n of Bone-Implan t Attachmen t Strengt h Between the Implant s With Hydroxyapatite-Coatin g and Tricalciumphosphate-Coatin g on Titaniu m Arc Spraye d Titaniu m

291

K. Hayashi, T. Hara, T. Imamura and Y. Iwamoto Effect of Hydroxyapatit e Coatin g on Bony Ingrowt h into Groove d Titaniu m Implant s

295

K. Hayashi, T. Mashima, K. Uenoyama, T. Hara and Y. Iwamoto Mechanis m of the Inflammator y Reactio n of Conventiona l Calciu m Phosphat e Cemen t

301

K. Ishikawa, Y. Miyamoto, M. Nagayama and K. Suzuki Fractur e of Alumin a Cerami c Head in Tota l Hip Arthroplasty . - Repor t of a Case With Histologica l Examinatio n and Particl e Characterisatio n

305

Y. Kadoya, A. Kobayashi, P.A. Revell, H. Ohashi, Y. Yamano, G. Scott and M.A.R. Freeman Experimenta l Comparativ e Stud y Between Rough-Blaste d and Hydroxyapatit e Coate d Implant s

309

Y.H. Kim, J.S. Park, I.Y. Choi, M.R . Park and T.S. Park Mechanica l and Biological Propertie s of Alumin a Bead Composit e

313

M. Kobayashi, T. Nakamura, T. Kikutani, Y. Okada, N. Ikeda, S. Shinzato and T. Kokubo Remodelin g of Bone Aroun d Hydroxylapatite-Coate d Femora l Stem s

317

A.A. Edidin and M.T. Manley Processin g and Characterisatio n of Biological Hydroxyapatit e Derived from Cattle , Sheep and Deer Bone

321

M.R . Mucalo , G.S. Johnson and M.A. Lorie r Catastrophi c Wear of Meta l Ball of Bipolar Hip Prosthesi s After Fractur e of Alumin a Cerami c Screws Used for Acetabula r Bone Graf t

325

H. Ohashi, Y. Yutani, A. Kobayashi, Y. Kadoya and Y. Yamano Antibacteria l Propert y of Ag-Doped Calciu m Phosphat e Compound-Cellulos e Composite s K. Okada, Y. Yokogawa, T. Kameyama, K. Kato, Y. Kawamoto, K. Nishizawa, F. Nagata, M. Okuyama

329

Contents xvii

Wear Behaviour of Polyethylen e Cup Against 28 mm Alumina Ball in Tota l Hip Prosthese s H. Oonishi, N. Murata, S. Kushitani, S. Wakitani, K. Imoto, Y. Iwaki and N. Kin In Vitr o Cell Behavior of Osteoblast s on Pyros t Bone Substitut e J-S. Sun, F.H. Lin, Y-H. Tsuang, Y-S. Hang, C.Y. Hong and H.C. Liu The Efficacy of Hydroxyapatite-Tricalciu m Phosphat e Filler for Bone Defects Associated With Humera l Pseudoarthrosis : Compariso n With Autogenou s Iliac Bone Graft s K. Suzuki and M. Yamada

333

337

341

Experimenta l Study of Apatit e Cement Includin g Cisplati n Y. Tahara, Y. Ishii, S. Sasaki, I. Takano and K. Ohzeki

345

In Vivo Evaluatio n of Sol-Gel Bioglassfi. - Biomechanica l Finding s D.L. Wheeler, R.G. Hoellrich, S.W. McLoughlin, D.L. Chamberland and K.E. Stokes

349

Fixatio n of Hip Prosthese s by Hydroxyapatit e Coatin g G. Willmann

353

Acetabula r Reconstructio n With an Artificia l Bone Block S. Yoshii, M. Oka, T. Yamamuro, H. lida, Y. Kakutani, K. Ikeda, H. Murakami and T. Nakamura

357

Participatio n of Calcium Phosphat e Cement s for Healin g of Alveolar Bone M. Yoshikawa, H. Oonishi, Y. Mandai, K. Minamigawa and T. Toda

361

BIOCERAMIC S SYNTHESI S AND EVALUATIO N Synthese s of Rapid Resorbabl e Calcium Phosphat e Ceramic s With High Macr o or High Micr o Porosit y G. Berger, R. Gildenhaar, U. Ploska and M. Willfahrt

367

Composit e Bioceramic s Mad e of Macroporou s Calcium Phosphat e Ceramic s FiUed With a Self-Settin g Cement . Histologica l Evaluatio n P. Frayssinet, A. Lerch, L. Gineste and N. Rouquet

371

xviii

Contents

Physica l Propertie s of an Apatiti c Cerami c Containin g Tricalciu m Phosphat e Prepare d by the Way of a Cement Z. Hatim, M. Freche and J.L. Lacout Cytocompatibilit y of Calcium Phosphat e Coating s With Variou s Ca/P Ratio s P. Frayssinet, L. Arbore and N. Rouquet Compariso n of Resorptio n and Bone Conductio n of Two CaCO a Bone Substitute s J.C. Fricain, Ch. Baquey, B. Basse-Cathalinat and B. Dupuy Reliabilit y of Dual Energ y X-Ray Absorptiometr y in Evaluatio n of Phospho Calcic Bioceramic s in Rabbi t J.X. Lu, O. Legrand, B. Flautre, A. Gallur, M. Descamps, B. Thierry, P. Hardouin and B. Sutter The Evaluatio n of Degradabilit y of Melt and Sol-Gel Derived Bioglassfi InVitr o D.C. Greenspan, J.P. Zhong, X.F. Chen and G.P. LaTorre

375

379

383

387

391

COMPOSIT E CERAMIC S Upgradin g of Hydroxyapatit e Cerami c Biocompatibilit y by Incorporatio n of a-Tricalciu m Phosphat e S. Sarig, F. Apfelbaum and F. Kahana

397

Preparatio n of Composit e Material s Calcium Hydroxyapatite/CoUage n by Coprecipitatio n Metho d O.I. SHvka and V.P. Orlovskii

401

Bony Reactio n of Severa l Kinds of Ca-P-Collage n Conjugate d Sponges H. Oonishi, F. Sugihara, K. Minamigawa, Y. Mandai, K. Nagatomi, S. Kushitani, H. Iwaki, N. Kin, E. Tsuji

403

In Vitr o and In Vivo Tests of Newly Developed TCP/CPL A Composite s M. Kikuchi, S-B. Cho, Y. Suetsugu, J. Tanaka, T. Kobayashi, M. Akao, Y. Koyama and K. Takakuda

407

Occlusion of Dentin Tubule s by 45S5 Bioglassfi L.J. Litkowski, G.D. Hack, H.B. Sheaffer and D.C. Greenspan

411

Contents xix

Bioiner t and Biodegradabl e Polymeri c Matri x Composite s Filled With Bioactive Si02-3CaO-P205-Mg O Glasses and Glass-Ceramic s R.L. Reis, A.M. Cunha, M.H. Ferdandez and R.N. Correia

415

The Healin g of Segmenta l Bone Defects, Induce d by Bioresorbabl e Calcium Phosphat e Cement Combine d With rhBMP- 2 ; Using as Past e K. Ohura, C. Hamanishi, S. Tanaka and N. Matsuda

419

DENTA L AND E.N.T . APPLICATION S Implan t Placemen t Enhance d by a New Bioactive Materia l E. Schepers and L. Barbier

425

Behaviour of Bioactive Glass (S53P4) in Huma n Fronta l Sinus Obliteratio n K. Aitasalo, J. Suonpaa, M. Peltola and A. Yli-Urpo

429

All-Cerami c Denta l Bridge s by the Direct Cerami c Machinin g Proces s (DCM) F. Filser, P. Kocher, H. Liithy, P. Scharer and L. Gauckler

433

Grindin g of Zirconi a - TZP in Dentistr y - CAD/CAM-Technolog y for the Manufacturin g of Fixed Dentures R. Luthardt , W. Rieger and R. Musil Zirconi a Implant s With a Plasma-Spraye d SiOi -HA Bioactive Coatin g A. Pedra and P. Sharroc k

437

445

ORTHOPAEDI C APPLICATION S Effect of Time and Temperatur e on the Productio n of Porou s Electrolyti c Hydroxyapatit e Coating s N. Asaoka, S. Best and W. Bonfield

447

Calcium Phosphat e Formatio n on Chemicall y Treate d Vacuum Plasm a Spraye d Titaniu m Coating s S.-W. Ha, K-L. Eckert , H. Gruner and E. Wintermante l

451

Propertie s of Plasm a Spraye d Bioactive Fluorhydroxyapatit e Coating s X. Ranz, C. Rey, N. Antolotti, M.F. Harmand, A. Moroni, L. Orienti, G. Viola, S. Bertini, A. Scrivani

455

XX

Contents

Longer-Ter m Mechanica l and Biological Evaluatio n of Titaniu m Alloy Coate d With Apatit e Layer

459

W.Q. Yan, K. Kawanabe, T. Nakamura and T. Kokubo Electrophoreti c Coating s of Porou s Apatit e Composit e onto Alumin a Ceramic s

463

K. Yamashita, E. Yonehara, J-i. Hamagami and T. Umegaki Osseointegratio n in Experimenta l HA-Coate d Femora l Stems

467

E. De Santis, G. Rinonapoli, C. Doria, A. Manunta and M.C. Sbernardori Effect of Hydroxyapatite-Coatin g on the Bondin g of Bone to Titaniu m Implant s in the Femu r of Ovariectomize d Rat s

471

T. Hara, K. Hayashi, Y. Nakashima and Y. Iwamoto

BI O ACTIV E BON E CEMEN T Optimizatio n of Settin g Time and Mechanica l Strengt h of ^-TCP/MCP M Biocement s

477

P. Van Landuyt, C. Lowe and J. Lemaitre Influenc e of the Particl e Size of the Powder Phas e in the Settin g and Hardenin g Behaviou r of a Calciu m Phosphat e Cemen t

481

M.P. Ginebra, E. Fernandez, F.C.M. Driessens, M.G. Boltong and J.A. Planell Subcutaneou s Tissue Response s and Kinetic s of Cells to Tetracalciu m Phosphat e Cement s

485

M. Yoshikawa, H. Oonishi, Y. Mandai, F. Sugihar and T. Toda Biological Behaviou r of a Bioactive Bone Cemen t Implante d in Rabbi t Tibia e

489

A. Afonso, M. Vasconcelos, R. Branco and J. Cavalheiro Histologica l Stud y of a DCPD-Base d Calciu m Phosphat e Cemen t

493

P. Frayssinet, L. Gineste, P. Conte, J. Fages, N. Rouquet and A. Lerch Bioactive Bone Cemen t Studie d in Canin e Tota l Hip Arthroplasty , 2 Year s FoUow-Up Stud y H. Fujita, T. Nakamura, K. Ido, Y. Matsuda, H. lida, M. Kobayashi, M. Oka and Y. Kitamura

497

Contents xxi

NE W MATERIAL S AND TECHNOLOGIE S

Experimenta l Stud y on Hydroxyapatite/N-Carboxymethy l Chitosa n Filler s

503

R. Martinetti, L. Dolcini, A. Ravaglioli, A. Krajevski and C. Mangano

Injectabl e Chitosamin e Hydroxylapatit e Bone Past e

507

J.J . Railhac, P. Sharrock, D. Galy-Fourcade, C. Zahraoui and N. Sans

Manufactur e of a Hydroxyapatite-Chiti n Composit e

511

A.C.A. Wan, E. Khor and G.W. Hastings

Load-Bearin g and Ductil e Hydroxylapatite/Polyethylen e Composite s for Bone Replacemen t

515

R.L. Reis, A.M. Cunha and M.J. Bevis

In Vitr o Assessment of Hydroxyapatite - and Bioglassfi - Reinforce d Polyethylen e Composite s

519

J. Huang, L. Di Silvio, M. Wang, K.E. Tanner and W. Bonfield

Osteoconductiv e Propertie s of Pur e and Type-A Carbonate d Hydroxyapatite s

523

S.A. Redey, D. Bernache-Assollant, C. Rey, P.J. Marie, M. Nardin and L. Sedel

Coagulatio n Times of Blood in Contac t With Gel-Derive d Silica-Alumin a Composit e Powder s

527

S. Takashima, C. Ohtsuki, S. Hayakawa and A. Osaka

Preparatio n of P"*"-Implante d Y203-Al203-Si02 Glas s for Radiotherap y of Cance r

531

M. Kawashita, F. Miyaji, T. Kokubo, G.H. Takaoka, I. Yamada, I. Suzuki and M. Inoue

New Ferromagneti c Bone Cemen t for Local Hyperthermi a K. Takegami, T. Sano, H. Wakabashi, J. Sonoda, T. Yamazaki, S. Morita, T. Shibuya and A. Uchida

535

xxii

Contents

BIOCERAMIC S PROCESSIN G Adsorptio n of L-Lysin e onto Silica Glass : A Synergisti c Approac h Combinin g Molecula r Modelin g With Experimenta l Analysi s

541

R.A. Latour Jr., J.K. West, L.L. Hench, S.D. Trembley, Y. Tian, G.C. Lickfield and A.P. Wheeler Effects of Divalent Cation s on Calciu m Phosphate s Precipitatio n on a Langmuir-Blodgett e Monolaye r

545

S.B. Cho, Y. Suetsugu, J. Tanaka, R. Azumi and M. Matsumoto Effect of Processin g on the Characteristic s of a 20 Vol.% AI2O3 Platelet Reinforce d Hydroxyapatit e Composit e

549

S. Gautier, E. Champion, D. Bernache-Assollant Mechanica l Evaluatio n of Phosphat e Biodegradabl e Glasse s by Mean s of Indentatio n Method s

553

F J . Gil, R. Terradas, J. Clement, G. Avila, S. Martinez and J.A. Planell Silicon in Connectiv e Tissue: Semi-Empirica l Molecula r Orbita l Model s

557

K.D. Lobel, J.K. West and L.L. Hench The Effect of Hea t Treatmen t on Bone Bondin g Ability of Alkali-Treate d Titaniu m

561

S. Nishiguchi, T. Nakamura, M. Kobayashi, W-Q. Yan, H-M. Kim, F. Miyaji and T. Kokubo Therma l Processin g of Compac t Bovine Bone

565

G. Vargas, M. Mendez, J. Mendez and J. Lopez

EVALUATION S M E T H O D S AND NE W APPLICATION S Characterizatio n of Syntheti c and Biological Calciu m Phosphat e Material s by Micro-Rama n Spectrometr y

571

G. Penel, G. Leroy, G. Cournot and E. Bres Biological Evaluatio n of Glas s Reinforce d Hydroxyapatit e by Flow Cytometr y M.A. Lopes, J.C. Knowles, K.A. Hing, J.D. Santos, F.J. Monteiro and L Olsen

575

Contentsxxiii

Evaluatio n of Macrophag e Respons e to Cerami c Particle s by Flow Cytometry : Analysis of Phagocytosi s and Cytotoxicit y

579

I. Catelas, R. Marchand, L’H. Yahia and O.L. Huk Stud y of Porou s Interconnection s of Biocerami c on Cellula r Rehabitatio n In Vitr o and In Vivo

583

J.X. Lu, B. Flautre, K. Anselme, A. Gallur, M. Descamps, B. Thierry and P. Hardouin Repai r of Osteochondra l Defect Using Artificia l Articula r Cartilag e

587

M. Hasegawa, A. Sudo, Y. Shikinami and A. Uchida Calciu m Phosphat e Ceramic s as Controlle d Release System s for FGF- 2

591

V. Midy , E. Hollande , C. Key and M. Dar d C-SRC Oncogen e mRNA Expressio n in Porou s Hydroxyapatit e Ceramic s K. Mishima, H. Ohgushi, T. Yoshikawa, H. Nakajima, E. Yamada, S. Tabata, Y. Dohi and K. Ichijima Autho r Inde x Keywor d Inde x

595


CALCIUM PHOSPHATE IN VIVO FORMATION


Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd

PRECIPITATIO N OF CALCIU M PHOSPHAT E ON TITANI A CERAMIC S K.-L. Eckert, S.-W. Ha, S. Ritter, E. Wintermantel Chair of Biocompatible Materials Science and Engineering, Department of Materials, ETH Zurich, Wagistrasse 23, CH-8952 Schlieren, Switzerland

ABSTRAC T Titania ceramics were prepared, pretreated with lOM NaOH solution and immersed in simulated body fluid (SBF) for up to 10 days. Morphological and chemical changes were analysed by SEM and EDX. NaOH-pretreatment lead to surface roughening of the titania grains and to formation of needle-like crystals. EDX analysis showed that Na was present at the surface. After immersion in SBF precipitation of spherical agglomerates occurred together with formation of a layer of needle› like crystals. Na in the surface disappeared and Ca became distinct after 10 days of immersion, suggesting a precipitation of calcium phosphate. Thus the reactions which were observed on titania surfaces after NaOH treatment are similar to those observed on titanium metal after identical treatment. KEYWORD S Titania ceramics, NaOH treatment, calcium phosphate INTRODUCTIO N The formation of biomimetic calcium phosphate layers on titanium surfaces immersed in simulated body fluid (SBF) was shown to occur after chemical treatment with alkaline solutions [1,2]. Due to the similarity of surface oxide layers on titanium metal and previously developed titania ceramics [3], it is assumed that the same procedure could also induce the formation of calcium phosphate layers on titania ceramics. The aim of the present study was to prove latter hypothesis that titania ceramics can be modified by sodium hydroxide (NaOH) treatment in a way that afterwards a calcium phosphate layer is formed if they are immersed in simulated body fluid. MATERIAL S AND METHOD S Ceramic Processing Titania ceramic discs of 10 mm diameter were prepared from titania powders by mixing 20 g of processed titanium dioxide powder (dso" 6.1 |Lim) with 12.0 g titania powder 1171 (Kronos, Germany), 1.0 g graphite powder KS6 (Lonza, Switzerland) and 2.8 g paraffin (MP 64 C, Fluka, Switzerland) into a thermoplastic body. The mixture was heated to 100 C and pressed into round discs of 15 mm diameter. For polymer burnout the samples were placed on an alumina refractory plate and heated to 300 C at a rate of 5 K/h. Final sintering was performed at 1350 C

4

Bioceramics Volume10

with a heating rate of 3 K/min and 25 min holding time. Finishing was done by ultrasonic cleaning for 15 seconds and flushing with deionized water. NaOH Treatmentand Immersionin SBF Alkaline treatment of the titania samples was carried out in lOM NaOH. The specimens were placed into a conical flask filled with 100 ml NaOH. Immersion was performed at 60 C for 2 hours in a laboratory shaker rotating at a speed of 70 rpm. After NaOH treatment the samples were gently flushed with deionized water for 1 minute. Immediately after soaking, the NaOH treated samples were placed into polypropylene vessels containing 25 ml simulated body fluid (SBF) prepared according to [5]. The pH of SBF was 7.4, pH control was performed at the end of every immersion period. The vessels were sealed and immersion in SBF was carried out at 37 C for 1, 4 and 10 days in a laboratory shaker rotating at 70 rpm. After the immersion in SBF the samples were thoroughly rinsed with deionized water and dried in ambient atmosphere at room temperature. Morphological and Chemical Characterization Morphology of the NaOH treated and of the immersed titania surfaces was analysed using scanning electron microscopy (SEM, Hitachi S-2500C). They were compared with control samples, not treated with NaOH, but immersed in SBF. Energy dispersive X-ray (EDX) analysis was performed at an acceleration voltage of 25 kV with an X-ray microanalysis system attached to the SEM (Voyager, Noran Instruments). EDX spectra were acquired with an acquisition time cf 100 seconds. The specimens were coated with platinum in a sputter coater prior to SEM and EDX analysis. RESULT S AND DISCUSSIO N NaOH Treatment Scanning electron microscopic evaluation of the surface of the untreated ceramic showed that the surface has a granular structure (figure 1) with growth steps on the individual grains which occur due to the crystalline nature of the material (figure 2). NaOH treated titania surfaces were similar to untreated samples with the difference that, at higher magnification, surface roughening on the titania crystallites and formation of needle-like structures was observed (figure 4). EDX analysis

Figure 1: Survey of untreated titania ceramic. The surface topography is defined by the granular structure of the material. The total porosity of the material is 25 %.

Figure 2: Untreated titania ceramic. The grains are structured by growth steps.

Precipitationof Calcium Phosphateon Titania Ceramics: K-L. Eckert et al.

5

(figure 7) of the untreated controls showed no other signals than titanium. After NaOH treatment Na was present at the surface. It is assumed that the newly formed, needle-like structures, which were observed with SEM (figure 4) contain sodium. Besides Na and Ti, no additional elements were found at the surface of NaOH treated specimens. Immersion in SBF After one day of immersion, the formation of globular precipitates and a fine layer of needle-like crystals were observed. The appearance of the precipitated layer did not change significantly during the following periods of immersion. After 10 days, the titania surface was completely covered with a layer consisting of needle-like crystals and spherulitic precipitations (figure 5). The thickness of the deposited layer as well as the amount of spherulites (figure 6) was markedly

I igure 3: I ilania ceramic after NaOHtreatment. No change in the topographical characteristics can be noticed.

Figure 4: Titania ceramic after NaOHtreatment. The surface was roughened, growth steps were partly etched away. Newly formed crystals occurred at the surface and in pores.

Figure 5: NaOH-pretreatment and 10 days of immersion in SBF. The surface is completely covered by a precipitated layer.

Figure 6: NaOH-pretreatment and 10 days of immersion in SBF. The precipitated layer consists of needle-like crystals and of spherical agglomerates.

6


c

3 O

O

10 Energ y [keV]

Figure 7: EDX spectra of the sample surfaces at various processing stages. On untreated samples only titanium signals could be detected. Pt signals are caused by the platinum sputter coating of the samples. After immersion in NaOH solution (Ti02+NaOH), a Na peak occurs. After 10 days of immersion in SBF, Ca is present, indicating the precipitation of calcium phosphate at the surface. increased compared to day one. However, the original topography of the titania ceramic surface was still maintained. In the EDX spectra, Ca was identified after the first day of immersion and became more distinct with time (figure 7). In contrast, Na vanished, probably due to ion exchange processes [5]. On control samples which were not treated with NaOH, no precipitation was observed. CONCLUSIO N The current work showed that precipitation of calcium phosphates on titania ceramics occurs similarly to those on titanium metal. In conclusion, the investigation has shown the potential of biomimetic calcium phosphate deposition on titania ceramics after pretreatment with lOM NaOH solution at 60 C for 2 hours. REFERENCES 1. Kokubo, T., Thermochimica Acta, 1996, 280/281, 479-490 2. Kim, H.-M., Miyaji, F., Kokubo, T., Nakamura, T., Journal of Biomedical Materials Research, 1996, 32, 409-417 3. Blum, J., Eckert, K.-L., Schroeder, A., Petitmermet, M., Ha, S.-W. and Wintermantel, E. In: Bioceramics Volume 9, Elsevier Science Ltd., Oxford 1996, 89-92 4. Kokubo, T., Hata, K., Nakamura, T., Yamamuro, T. In: Bioceramics Volume 4, Butterworth-Heinemann, Guildford (1991), 113-120 5. Clearfield, A., Lehto, J., Journal of Solid State Chemistry, 73 (1988), 98-106


APATIT E FORMATIO N ON POLYMER S BY BIOMIMETI C PROCESS USING SODIU M SILICAT E SOLUTIO N F. Miyaji^, S. Handa\ T. Kokubo^ and T. Nakamura^ ^ Department of Material Chemistry, Faculty of Engineering, Kyoto University, Yoshida-honmachi, Sakyo-ku, Kyoto 606-01, Japan ^Department of Orthopaedic Surgery, Faculty of Medicine, Kyoto University, ShogoinKawaharacho, Sakyo-ku, Kyoto 606-01, Japan

ABSTRAC T A variation of biomimetic process which aims at a bonehke apatite coating on organic polymers with complex shapes was attempted by using sodium silicate as a catalyst for the apatite nucleation, and a simulated body fluid as a medium for the apatite growth. An apatite-forming abihty was the highest when the Si concentration and Si02/Na2 0 ratio of the sodium siUcate solution were above 2.0 M and 1.0-1.5, respectively. Particular sihcate ohgomers were assumed to be most responsible for the apatite nucleation. The apatite layer was formed not only on the flat PET surfaces but also on curved surfaces of fine PET fibers constituting a fabric, where the apatite layer was interconnected each other. This method is expected to enable the bonelike apatite coating on various kinds of materials with complex shapes. KEYWORDS : Apatite, Biomimetic process, Sodium silicate, Simulated body fluid INTRODUCTIO N A biomimetic process has been developed for coating a dense and uniform bonelike apatite layer on organic polymers as follows [1,2]. First, in order to form apatite nuclei on the substrates of organic polymers, the polymer substrates are placed on CaO-Si02- or Na2 0-Si02based glass particles soaked in a simulated body fluid (SBF) with ion concentrations nearly equal to those of human blood plasma [3]. Next, in order to make the apatite nuclei grow, the polymer substrates are soaked in 1.5SBF with ion concentrations 1.5 times the SBF. The disadvantage of the above biomimetic process lies in the diflficulty of apatite coating on the materials with complex shapes, since in the first treatment the apatite nuclei are formed only on the material surface which is faced to the glass grains. In the present study, the apatite formation on organic polymers was attempted by using sodium silicate solution as a nucleating agent for the apatite formation instead of the glass particles, and 1.5 SBF as a medium for the apatite growth. MATERIAL S AND METHOD S Preparatio n of sodiu m silicat e solutio n Reagent grade sodium metasilicate (Na2 Si03) was dissolved into distilled water to prepare solutions with 0.5, 1.0, 2.0 and 3.0 M Si-concentration. As another series of sodium silicate

8

Bioceramies Volume10

solutions, the solutions with SiO./Na.O ratio of 0.5, 0.67, 0.8, 1.0, 1.5 and 2.0 were prepared by adding reagent grade NaOH or SiO ^ xH^ O into sodium metasihcate solution, where Si concentration was fixed at 3.0 M. Preparatio n of l.SSBF The 1.5SBF with ion concentrations (Na^ 213.0, K^ 7.5, Mg^^ 2.3, Ca^^ 3.8, CI’ 223.2, HCO3’ 6.3, HPO/ 1.5, s o / 0.8 mM) 1.5 times the SBF was prepared by dissolving reagent grade NaCl, NaHC03, KCl, K^HPO.^HÔ, MgCl^^HÔ, CaCl^ and Na^SO^ into distilled water. The pH of the solution was adjusted at 7.25 with NH^QCHÔH) and 1 M-HCl at 36.5T. Apatit e coatin g on polyme r Rectangular substrates (10 x 10 x 1 mm ) of poly (ethylene terephthalate) (PET) were abraded #400 and washed with ethanol. And then the substrates were subjected to a glow discharge treatment in O2 gas for 30 s [4] for producing polar groups on the polymer surfaces, which might contribute to the strong attachment between silicate ions and the substrate. After the treatment, the substrates were soaked in sodium sihcate solutions with various concentrations and Si02/Na20 ratios for 6 h at 36.5T. After removing from the solution, the substrates were dried at room temperature, rinsed with distilled water and then soaked in 20 ml of 1.5 SBF for various periods. A fabric (10 x 15 mm^ in area) woven with ultrafme PET fiber (2 \xm^)(Toray Co. Ltd., Otsu, Japan) was also used as a substrate. Surfac e analysi s After the soaking in 1.5 SBF, surface structural and morphological variations of the specimens were characterized by a thin-film X-ray diffractometer (TF-XRD; thin-film attachment CN2651A1, Rigaku-Denki Co., Tokyo, Japan), a Fourier transformed infrared (FT-IR) reflection spectrometer (System 2000 FT-IR, Perkin-Elmer Ltd., Buckinghamshire, England) and a scanning electron microscope (SEM; S-2500CX, Hitachi Co., Tokyo, Japan). RESULT S AND DISCUSSIO N Effect of concentratio n of sodiu m silicat e solutio n Figure 1(a) shows the TF-XRD patterns of the surfaces of PET substrates soaked in 1.5 SBF for 6 d after soaking in sodium metasihcate solutions with various concentrations for 6 h. The peaks ascribed to apatite were observed for the 2.0 and 3.0 M-treated specimens. Figure 2 shows the SEM photographs of the surfaces of PET substrates soaked in 1.5 SBF for 6 d after soaking in sodium metasihcate solutions with various concentrations for 6 h. In the case of 0.5 M, apatite particles were deposited only on a small part of the surface of the substrate. The number of apatite particles increased for the treatment with 1.0 M solution. Moreover, whole surfaces of the substrates were covered with apatite layer for the treatment with 2.0 and 3.0 M solutions. These results indicate that the apatite-forming tendency becomes higher with increasing concentration of sodium sihcate solution. This is explained by assuming that the number of sihcate ions attached to the substrates increased with increasing concentration of sodium silicate solution, forming more apatite nuclei. It should be, however, noted that the degree of the apatite formation is almost the same between 2.0 and 3.0 M. This suggests that the number of sihcate ions attached to the substrates are saturated at 2.0 M concentration. Effect of compositio n of sodiu m silicat e solutio n Figure 1(b) shows the TF-XRD patterns of the surfaces of PET substrates soaked in 1.5SBF for 6 d after soaking in sodium silicate solutions with various Si02/Na2 0 ratios for 6 h. The peaks ascribed to apatite were observed for 1.0 and 1.5 of Si02/Na20 ratios. Figure 3 shows the SEM photographs of the surfaces of PET substrates soaked in 1.5 SBF for 6 d after

Apatite Formation on Polymers by BiomimeticProcessing: F. Miyaji F. et al. 9 O Apatite

O Apatite

SiOg / Na20

20 30 40 50 60 i^ 20 30 40 50 60 26 (CuKa) / degree 29 (CuKa) / degree Figure 1 TF-XRD patterns of the surfaces of PET substrates soaked in 1.5SBF for 6 d after soaking in sodium silicate solutions with (a) various concentrations and (b) various Si02/Na20 ratios for 6 h. 10

fi 6 ' D3

go fi CD

mMm

0 fi O O O u D Rt

’0fiOQQuD

Figure 2 SEM photographs of the surfaces of PET substrates soaked in 1.5SBF for 6 d after soaking in sodium metasilicate solutions with various concentrations for 6 h.

Hofi

o'

iofi tlfiCQQuD

’’Wm

Figure 3 SEM photograplis of the surfaces of PET substrates soaked in 1.5SBF for 6 d after soaking in sodium sihcate solution with various Si02/Na2 0 ratios for 6 h.

10 Bioceramies Volume10

soaking in sodium silicate solution with various Si02/Na2 0 ratios for 6 h. No apatite was observed for 0.5 of ^\0^f^?iÔratio. The whole surfaces of the substrates were covered with apatite for 1.0 and 1.5 of Si02/Na20 ratios. Apatite particles were deposited on the substrates in part for 2.0 of Si02/Na2 0 ratio. These indicate that the sodium silicate solutions with 1.0 and 1.5 of Si02/Na2 0 ratios are most adequate for apatite formation. It is well known that the structure of sihcate ions changes with the composition of sodium silicate solutions [5]: silicate ions are primarily present as monomer for SiO2/Na2O tetraethoxysilane (TEOS), Si(OC2H5)4) and calcium nitrate (Ca(N03)2*4H20). One of the typical compositions for starting materials was: TEOS:PDMS:HCl:H20:Ca(N03)2*4H20 = 1 : 1.67 : 0.05 : 3 : 0.05. Detailed procedure was given in the previous reports[3,4]. In the present study, 33

34


the ratio of HCl and Ca(N03)2«4H20 was varied in the range of Ha/TEOS=0.05-0.1(mol) and Ca/TEOS=0’-0.1(mol), respectively. The synthesized samples were cut to 15x10x1 mm^ in size, and the surface was polished with a #2000 emery paper. Then they were gently rinsed with ethanol and dried. Microstructure of the fracture-surface was observed with a scaiming electron microscope(SEM), JEOL JSM-6300, equipped with an energy dispersive X-ray analyzer (EDX, PhiUips EDX-4). The specimens were soaked in a simulated body fluid(Kokubo solution) at 36.5 C up to 14 days. The Kokubo solution was prepared as described by Cho et al\5\. It has been proved[6] that the in vivoapatite formation on implants can almost fully be reproduced in vitroexperiments in the Kokubo solution. Thin-film X-ray diffraction(TF-XRD) was used to examine formation of hydroxyapatite layer on the surface of the specimens after soaked in the Kokubo solution. To evaluate local structure around Si atoms in the samples, ^^Si Cross-polarization Magic Angle Spiiming (CP MAS) NMR spectra was measured for the samples after pulverized. In the NMR measurement, the sample spiiming speed was 5-6kHz at the magic angle(54.7 ) to the external field. 29si NMR spectra of the samples were measured at 9.4T on a JEOL JNMGX400 FT-NMR spectrometer, equipped with a TU-GSX400 MAS probe. 29si NMR spectra were acquired at 79.3MHz with 5.0-(xs pulses, 12.0s recycle delays, 15.0(xs dead time. The signals for an 80’-400 pulses were accumulated. The chemical shift (6 in ppm) of 29si was determined using tetramethylsilane(TMS)(6=0 ppm : ^^Si chemical shifts where 6 denoted the isotropic chemical shift) as an external reference substance. Poly dimethyl silane(6=-34.0 ppm) was used as the secondary external reference. RESULT S AND DISCUSSIO N Figure 1 shows the TF-XRD patterns of the samples with varied ratio of HCl after soaked in the Kokubo solution up to 14 days. They are denoted as "HCl X Ca Y " where X is the molar ratio HCl/TEOS(mol) and Y is the molar ratio Ca/TEOS (mol). Fig. 1 indicates that apatite formation on the Ormosils-type samples with a fixed ratio of Ca(N03)2*4H20 was enhanced by the increasing of the amount of HCl added. For the samples with varied ratio of Ca(N03)2*4H20 at H Q of Y=0.1, apatite formation was enhanced by the increasing of the amount of calcium nitrate added. Thus the bioactivity of the Ormosils-type materials was effected not only additional amounts of Ca(N03)2*4H20 but also the amounts of H Q solution. The results of 29si CP MAS NMR measurements shows that two groups of Si atoms were distinguished on the basis (rf the chemical shift: the first peak(-10 to -25 ppm) was due to 02SiMe2(Me=-CH3) units from PDMS while the other peak(-100 to -110 ppm) was due to Si02 units originated from TEOS. Both peaks were very broad because they were the envelopes of a few component peaks. Deconvolution of them on the basis of the chemical shift data in the literature[7] indicates that the former peak had three components and the latter had two. The former curve was fitted with Lorentz functions and the latter was fit with Gaussian functions. Each component in the former peak is assigned to PDMS chain(D) , cyclic ohgomers(E)cycUc) and copolymerized species(D(Q)), respectively. The Si04 units for the latter peaks are denoted as (y^ groups where n represents the number of bridging oxygen atoms around a Si atom in the siUca matrices. Thus obtained peak positions and the relative peak area (%) are given in Table 1. Table 1 indicates that amount of H Q does not show a significant effect for the local structure of Si atoms. Similar was found for the effect of Ca(N03)2*4H20. Fig. 2 shows the results of SEMEDX analysis of fractured surfaces for Ormosils-type samples. Size of pores in the samples seems to increase with increasing the ratio of HCl solution. In addition, the concentration of Ca

Bioactivity and Structureof Organically Modified Silicate: K. Tsuru et al.

25

25

30 35 2e/de g

30 35 2e/de g

25

35

30 35 2e/de g

Figur e 1 TF-XR D pattern s of sample s synthesize d with variou s amount s of HCl after soaked in the Kokub o solution up to 14 days. O : Apatit e

in the sample s also increase d with increasin g the rati o of H Q and Ca(N03)2*4H20 . Therefore , greate r apatite-formin g abilit y has been attribute d to incorporatio n of mor e calcium ions in the polymer due to large r amount s of HQ . Dissolution of Ca(II ) ions in the structur e not only favor s formatio n of a sihca hydroge l layer conductiv e to apatit e nucleatio n but increase s in degre e of siq)ersaturatio n for âtit e in the Kokub o solution . Therefore , we conclude d tha t controllin g the incorporatio n of Ca(II ) ions is essentia l for preparatio n of bioactiv e Ormosils-typ e organicall y modified ceramics .

Tabl e 1. Peak position s and relativ e peak are a (%) of the 29si CP MAS NMR peaks . The left tabl e is due to 02SiMe2 units(Dgroup ) from PDMS. Theright tabl e is due to Si02 units(Q ^ group ) originate d from TEOS . Dg roups Sampl e nam e

0*^9 roups

6 (relative pea k area , % ) D(Q)

Dcydic

6 (relative pea k area , % )

D

Q3

Q4

HCl 0. 1 -CaO.OS

-17. 2 (28 )

-19. 6 (51 )

-21. 7 (21 )

-102. 3 (24% )

-108. 5 (76% )

HCl 0.0 9

C a 0.0 5

-17. 1 (29 )

-19. 7 (53 )

-21. 9 (18 )

-102. 7 (22 )

-108. 6 (78 )

HCl 0.0 7

C a 0.0 5

-16. 9 (30 )

-19. 7 (49 )

-22. 0 (21 )

-102. 9 (28 )

-108. 7 (72 )

HCl 0.0 5

C a 0.0 5

-17. 0 (34 )

-19. 6 (45 )

-21. 9 (21 )

-103. 0 (31 )

-108. 6 (69 )

36


HCI 0.0 5 rCa 0.0 5

SiKa 99.8% I CaKaO.2 %

SIKa 984%; CaKal6%

Figure 2 SEM-EDX analysis of fracture-surfaces for Ormosils-type samples. Apatite : X = no apatite after 14 days Apatite 3 days = apatite deposited after 3 days

CONCLUSIO N Apatite-forming ability of Ormosil-type organically modified ceramic is enhanced by controlled amounts of HCl and Ca(N03)2*4H20. Analysis of ^^Si MAS NMR spectra indicates that additional H Q does not show any significant effects for local structure around Si atoms. The porosity and Ca content incorporated in the structures depends on amounts of HCl. We have concluded that controlling the incorporation of Ca(II) ions is essential for preparation of bioactive Ormosils-type organically modified ceramics.

REFERENCES

1. Hench, L. L. and Wilson, J. In: An introductionto bioceramics.World Scientific, Singapore, 1993,1-24. 2. Ohtsuki, C , Kokubo, T. and Yamamuro, T., J. Non-Cryst.Solids, 1992,143,84-92. 3. Tsuru, K., Ohtsuki, C. and Osaka, A. 7. Mat. Sci.,Mat. Med., 1997,8, 157-161. 4. Tsuru, K., Hayakawa, S., Ohtsuki, C. and Osaka, A. In: BioceramicsVolimie 9, Pergamon, Oxford, 1996,419-422. 5. Cho, S.B., Nakanishi, K., Kokubo, T., Soga, N., Ohtsuki, C , Nakamura, T., Kitsugi, T. and Yamamuro, T. J. Am. Ceram.Soc, 1995, 78, 1769-1774. 6. Kokubo, T., Kushitani, H., Sakka, S., Kitsugi, T. and Yamamuro, T., J. Biomed.Mater. /?ej.,1990, 24,721-734. 7. Iwamoto, T., Morita, K., and Mackenzie, J. D.,J. Non-Cryst.Solids, 1993,159,65-72. ACKNOWLEDGEMENT S The authors thank Prof. T. Yoko of Kyoto University for his helpful advice and assistance in the NMR measurements. One of the authors(K. T.) gratefully acknowledges the Research Fellowship of the Japan Society for the Promotion of Science for Young Scientists.


HYDROXYAPATIT E FORMATIO N ON BIOACTIV E GLASS COATE D TITANIU M Cheol Y. Kim and Sungmin Kwon Dept. of Ceramic Engineering, Inha Univ. 253, Yonghyun-dong, Nam-ku, Inchon,402-751. Korea.

ABSTRAC T Two different bioactive glasses were coated on a titanium to give it a bioactivity, and then the bonding characteristics between a titanium and a coated layer and the hydroxyapatite formation in a simulated body fluid(SBF) on the coated specimens were studied. TisSis was found at the interface between titanium and glass coat in the groundcoat layer, and this enhaced the titanium-glass bonding. The hydrox}âpatite formation was observed only on the covercoat fired under 800 C, which is in an amorphous phase, and was retarded for the sample fired over 850 C, which contains an oxyapatite. KEYWORDS : Bioactive Glasses, Titanium, Hydroxyapatite, Groundcoat, Covercoat INTRODUCTIO N It has been well known that bioactive glasses are one of the best candidates for an implant material because of their excellent bonding behavior to the living tissue. ^^ The major drawback of these materials to use as a practical implant is their poor mechanical strength. To solve this problem, several works on coating the bioactive glasses to a strong substrate such as alumina and metal have been carried out.^^ A metal substrate has an advantage over an alumina because the former has better mechanical properties. However, most of metals easily oxidize specially at high temperarure and release some harmful ions when implanted for a long period of time. A titanium is now known as one of the best biometals for an implant material because it is chemically inert in a body. In the present study, therefore, two kinds of bioactive glasses are coated on a titanium by using a double coating method. The primary objectives of this study are 1) to find the bonding behavior between glass and titanium, and 2) to find the hydroxyapatite formation on the glass coated layer in a simulate body fluid. 37

38


MATERIAL S AND METHOD S Two types of glass powder, as shown in Table 1, were prepared by melting the glass batch in a Pt-crucible at 1500 C for 2 hours followed by pulverizing them to less than 44/mi in an agate mortar. The glass powder was dispersed in an aceton, and spray-coated on a titanium. Glass 55SF and 5 5 SB were used for a groundcoat and a covercoat, repectively. The glass coated titanium was fired in a tube furnace at the temperature ranging from 1150 Cto 1300 C. Ar gas was flowed into the tube furnace at the rate of 2.5 1/min to prevent the oxidation of the titanium. The bioactive glass coated titanium was reacted in a simulated body fluid (SBF), which has pH of 7.3, and the ratio of the bioactive glass coated area to volume of the reaction solution was set at 0.1 cm’^ and the reaction temperature at 37 C. The crystalline phases of the heat-treated bioactive glass coated layer and the hydroxyapatite crystalline phase after the reaction test were examined by a thin film x-ray diffractometer. Tabl e 1. Batch Compositions of Bioactive Glasses

(mol%)

sample

Si02

P2O5

Na.O

CaO

Cap.

B2O3

55SF

55.1

3.4

9.2

27.8

4.5

-

50SB

50.1

3.4

9.2

32.3

-

5

RESULT S AND DISCUSSIO N Bondin g Behavior between Glass and Titaniu m When 55SF glass was coated on a pure titanium and fired under Ar-gas condition at 1300 C for 2 minutes, the crystalline phases of the glass layer are shown in Fig.l. TisSia crystal was found at the interface, and this crystal promotes the bonding of the glass layer to the titanium. 55SF glass coated Ti-metal was fired between 1150 C and 1300 C, and the XRD patterns are shown in Fig.2. Afluorapatitecrystal was obtained at a lower firing temperature. Thefluorapatitewas mehed as increase in temperature and a -wollastonite was developed at 1300 C. It is believed that titanium ions penetrated into glass layer at a higher firing temperature and acted as a nucleating agent for an (7-wollastonite crystal formation. Generally an oxyapatite and fluorapatite crystal form at a lower temperature without nucleating agent. No hydrox>apatite formation was observed when these samples were reacted in the SBF.^^

Hydroxyapatite Formation on Bioactive Glass Coated Titanium:C. Y. Kim and S. Kwon 39

w a-Wollastonite

W a-Wollastonit e W

I

I

T Titanium

|i>ii MmA»nMû,0umiiK^

i3ao c

50^m

wl w s gs ^

I w

^

1250X

w

c

F Fluorapatit e

83p.m

100^m

I t I i I 20

30

40

50

26 (degree)

60

70

10

20

30

40

50

60

20 (degree)

Fig.l XRD patterns of 55SF glass coated Fig.2 XRD patterns for the surface of 5SF with increase in depth, heat-treated coated titanium, heat-treated at variat 1300 C in Ar atmosphere ous temperature for 2 mins Hydroxyapatit e Formatio n on Overcoate d Glass Layer Because no hydroxyapatite was formed on the groundcoat, another low melting glass 50SB was overcoated on the 55SF, and fired at 750T- 900T for 2 min. The glass layer crystallized into an oxyapatite when heat-treated at 900 C, and an amorphous overcoat layer was obtained when heat-trested at 800 C as shown in Fig. 3. These double-coat samples were reacted in SBF for 48 hours, and their XRD patterns are shown in Fig.4. The samples heat-treated at 850 C and 900 C showed oxyapatite crystal which were formed during the heat treatment. The sample heat-treated at 800’’C, however, showed a typical X-ray diffraction pattern of a newly formed hydroxyapatite precipitated on the glass surface. The reason for the delay of the hydroxyapatite formation on the oxyapatite crystal containing samples is because the matrix phase txu’ns into a chemically durable phase, and most of phosphorus ions are trapped in the oxyapatite crystal. This indicates that the leaching of phosphorus ions from the bioactive glass plays a very important role for the easy formation of the hydroxyapatite even in the phosphorus ion containing solution like SBF.

40

Bioceramics Volume10 0 Oxyapatite O Oxyapatite

1

H Hydroxyapatite 1

01

1 I L*M**«M(

900 C

KaJLjimlJliil

L-H’*»W\.

850 C 1

i " L^NIWWV K 1 ""^

20

30

40

50

60

20 (degree)

Fig.3 XRD patterns for the surface of 50SB coated titanium. Heat treated at various temperatures for 2 minutes.

10

800 C

:Tr^

1

J .1 .i_L _

10

1 1

20

30

40

50

60

2e (degree)

Fig.4 XRD patterns for the surface of 50SB coated titanium after corrosion in trisbuffer solution for 48 hrs.

CONCLUSIONS Bioactive groundcoat glass has a good chemical bonding to titanium by producing TisSis crystal at the interface. No hydroxyaptite formation was observed on the groundcoat after reaction in SBF, while hydroxyapatite crystal was obtained on the covercoat glass, which was fired at the temperature lower than 800 C. REFERENCE S 1. P. Ducheyne, P. Bianco, S. Radin and E. Schepers, Bone-BondingBiomatehals,Reed Healthcare, conmiun., 1992, 1-12 2. J. K. Kim and C. Y. Kim , J.Kor.Ceram.Soc, 1990, Vol.27, No.7, 925-933 3. C. Ohtasuki. A. Osaka and T. Kokubo, Bioceramics,1994, Vol. 7, 73-78


EFFEC T OF MULTIVALEN T CATION S IN CALCIU M SILICAT E GLASSE S ON BIOACTIVIT Y Naoki Imayoshi, Chikara Ohtsuki, Satoshi Hayakawa and Akiyoshi Osaka Biomaterials Lab., Faculty of Engineering, Okayama University, Tsushima, Okayama 700, Japan

ABSTRAC T We examined the apatite formation for CaOSiO^ glasses containing Vp^, WO3, Ta205or Cr03 after they were soaked in a simulated body fluid(SBF) for various periods. Surface reaction of glass with SBF and reconstruction of an Si-O-Si network in the surface layer were studied with ^^Si MAS-NMR spectroscopy which gave the average ratio of bridging/nonbridging oxygen for an Si atom(Q"). It was shown that TaÔ^ and Cr03 effectively depressed the apatite formation and that VÔg and WO3 up to 5mol% caused little influence while dissolution of Ca( II) and Si(IV) was enhanced. Analysis of the ^^Si MAS-NMR spectra showed that the distribution of Q" was different for each transition metal oxides, and thus suggested that the structure of silica hydrogel layer was affected by those cations. INTRODUCTIO N Bioactive glasses and glass-ceramics bond to living bone directly[l]. The condition for those to bond to living bone, i.e. bioactivity, is to form a bone-like apatite layer when embedded in the bony defect. A binary glass 50CaO-50SiO2(mol%) forms the apatite layer, hence it serves a basic system for bone substitutes. Addition of third components like AI2O3 and TiO, depresses the apatite formation [2]. Thus, those cations cause opposing effects on bioactivity of glasses as they modify chemical properties and structure of a silica gel layer that is assumed to favor the apatite nucleation. It is essential to understand what structural and chemical role the third cations have in bioactive glasses. Transition metal cations are worth examining such effects since their structural roles may depend on glass composition. In this study we examined the apatite formation of SOCaO-SOSiO^ glasses doped with 5mol% transition metal oxides as well as dopant-free glass soaking in a simulated body fluid(SBF) which are similar in inorganic composition to that of the human blood plasma. Moreover, ^^Si Magic Angle Spinning(MAS)-NMR spectra of pulverized glasses were measured before and after soaked in the SBF to determine glass structure. MATERIAL S AND METHOD S Glasses were prepared due to an ordinary melt-quench method in a series of composition: x (V2O3, WO3, Tap^ or CrO3)(50-x/2)CaO(50-x/2)SiO2 where the molar ratio CaO/SiO, was maintained to 1. The obtained glasses were cut into rectangular specimens 15 X 10 X Imm^ and polished with a diamond paste. They were soaked in the SBF with pH=7.25 at 36.5 C. The SBF was prepared as described by Cho eta/[31. After soaking the glasses into the SBF, their surface structure was examined by FT-IR spectroscopy and thinfilmX-ray diffractometry(TF-XRD). The concentrations 41

42


of calcium, silicon and phosphorus in the SBF after each period of soaking were measured by inductively coupled plasma(ICP) emission spectroscopy. ^S i MAS-NM R measuremen t The glanular glasses with 150-300îm in diameter were served for the^^Si MAS-NMR measurement before and after soaking in the SBF for various periods. ^^Si MAS-NMR spectra were recorded at 9.4T on a JEOL JNM-GSX400 FT-NMR spectrometer, equipped with a TU-GSX400MAS probe. Samples were placed in a zirconia sample tube. The sample spinning speed at the magic angle to external field was 5-6kHz. ^^Si MAS-NMR spectra were measured at 79.3MHz with 4.0-^s pulses, 2.5s recycle delays. About 1000 pulses were accumulated. Poly dimethyl silane(PDMS) was used as secondary external reference substance to determine the -^Si chemical shifts. RESULT S AND DISCUSSIO N Figure 1 shows the TF-XRD patterns of the surfaces of the metal oxide-doped and oxide-free 50CaO-50SiO, glasses soaked in the SBF for 7 and 30 days. The diffraction peaks near 26 and 32 were assigned to (002) and an envelope of (211), (112) and (300) of apatite, respectively. Fig. 1 indicates that SOCaO-SOSiO^ glass, 5mol%V203 and 5mol%W03 doped-glasses formed apatite on their surfaces in the SBF within 7 days, whereas 5mol%Ta205 and 5mol%Cr03 doped-glasses did not form it even after 30 days. This difference on apatite-forming ability among glasses suggests changes of chemical property and structure of SOCaO-SOSiO^ glass due to the addition of the transition metal oxides. The concentrations of calcium and silicon of the SBF due to soaking the glasses are shown in Figs. 2 and 3. They increased with the period for 5mol%V205 and WO3 doped-glasses, whereas only a slight

(b)30d Q

X-ra y 50CaO’50SiO 2

’’^w V ^

5mol%V,O s

0

5mol%W0 3

CI

e

A 2 _AÂ^ _ 5mol%Ta20 5 5mol%Cr0 3 ^ *f*l’*rrHrT i 1 1 1 1 1 Trt’^’i"^"’* "

20

25

30 35 2e/de g

40

20

25

30 35 2e/de g

40

Figure 1 TF -XRD patterns of 50CaO-50SiO 5mol% Vp^, WO3, Ta203 and Cr03 doped-glasses soaked in the simulated body fluid for 7 days (a) and 30days (b). O: Apatite

Effects of MultivalentCations in Calcium Silicate Glasses on Bioactivity:N. Imayoshi et al. 10 §2. 5 S ^2. 0 .0

(b)[SJ ]

-J

_

J

^,^

1 1.5

ii.o

-

s o

O.Or

10

15 20 Time/da y

25

43

-^ crS^m 15 20 i 25

10

6

30

135

Tim e / da y

Figure 2 Changes in Ca (a) and Si (b) concentrations of the simulated body fluid due to immersion of 50CaO-50SiO 5mol% V p ^ , WO3, Idifi^ and Cr03 doped-glasses in the simulated body fluid. D: 50CaO-50Sia, A: 5mol%V203, A: 5mol%W03, O: 5mol%Tap3 and FI: 5mol%Cr03. "

increase was noticed for 5mol% Cr03 doped-glasses. There appears no changes in the concentrations for 5mol% Ta203 -doped glass. This means the addition of TaÔ, improves chemical durability of 50CaO-50SiO^ glass and suppresses the dissolution of calcium and silicon. The increase in calcium favors the apatite formation due to increase in the degree of supersaturation with respect to hydroxyapatite, and the release of silicon from glasses shows formation of silanol groups which plays an important role on apatite nucleation. Similar analysis indicated release of vanadium, chromium and tungsten from the glasses in the SBF, although tantalum ion was not detected even after soaking for 30 days. Therefore, the dissolution of vanadium and tungsten ions from their glasses in the SBF did not affect the apatite formation. In contrast, 5mol%Cr03 doped-glasses did not form the apatite layer although release of calcium and silicon was detected. These results suggest that the dissolved chromium ions in the SBF remarkably suppresses the nucleation and/or growth of the apatite. Figure 3 shows ^^Si MAS-NMR spectra of 50CaO-50SiO,, 5mol%V,03, WO3, Ta,05 and Cr03 doped-glasses before soaking in the SBF. Each peaks were deconvoluted into two or three Gaussian functions on the basis of least square fitting. Each component was assigned to Q^ ( ? or Qôn the basis of the reference data[4], where Q" shows the number of bridging oxygen. Relative peak area of each Q" group (n=2-4) of 50CaO-50SiO, does not differ from that for 5mol% transition metal oxides. This means that the apatite formation is not associated the initial local structure around an Si atom. Fig. 4 shows 2^Si MAS-NMR spectra of 5mol% Vpând WO3 doped-glasses after soaking in the SBF for various periods. Leaching of calcium ions from the glass surfaces forms the Si-OH bonds of Q2 and Q^ groups without changing much of their ratios. However, the fraction of Q^ decreased for 5mol%V205 and WO3 doped-glasses with longer periods, probably due to polymerization of the SiOH bonds of Q2 groups. The increase in the fraction of Qt is attributed to the formation of Q^ units on the glass surface in the SBF. From Figs. 1 and 4, it is concluded that Q"^ units are favorable to apatite formation. Similar analysis for 5mol%Ta205 and Cr03 doped-glasses suggested absence of such changes in Q^ and Q^ for those glasses. A high field chemical shift, found in Fig. 4, for Si in the Q^ groups suggests that some of the Q^ groups involve Si-O-P bonds.

44

Bioceramics Volume10 I I H | H I H I I I I | I I I I | I I I 1 | I I I I | I I H |IH I I | I I I

’S i MAS-NM R spectr a

I l l i | n i l 11I I i | I II I I n i l | i i i 1 1 I I 1 1 | i 1 1 1 1 I I I

5mol % VÔ g

Q-

^4

50CaO-50SiO ,

5mol%Cr0 3 tHliliiiilmiLifH l

Figure 3 ^^Si MAS-NMR spectra of pulverised 50CaO-50SiO,, 5mol%V,03, WO3, Ta,03 and CrOj doped-glasses before soaking in the simulated body fluid.

Figure 4 ^^Si MAS-NMR spectra of pulverised 5mol%V20g and WO3 doped-glasses after soaking in the simulated body fluid.

SUMMAR Y The apatite formation for 50CaO-50SiO, glass respectively varied by addition of transition metal oxides. The addition of 5mol% VÔ^ and WO3 hardly effected the apatite formation in the SBF, while the addition of 5mol% TaPg and CrOj remarkably suppressed the apatite formation. The addition of the former increased in calcium and silicon concentration and the structure of silica gel layer as a result of chemical reaction in the SBF favored the apatite formation. But the addition of the latter improved the chemical durability and suppressed the formation of silica gel layer. Moreover dissolution of chromium ion in the SBF suppressed the apatite formation. REFERENCES 1. Hench, L.. L., Splinter, R. J., Allen, W. C. and Greenlee, T K. J, Biomed.Mater.Res. Symp.,2, 117-141(1972). 2. Ohtsuki, C , T. Kokubo, T, Takatsuka and T. Yamamuro, J. Mater.Sci.: Mater Med.,3, 119125 (1992). 3. Cho, S.B., Nakanishi, K., Kokubo, T, Soga, N., Ohtsuki, C , Nakamura, T, and Yamamuro, T. J. Am. Ceram.Soc, 78, 1769-1774 (1995). 4. Selvaray, U., Rao, K. J., Rao, C. N. R., Klinowski, J. and Thomas, J. M. Chem.Phys. Lett.,114, 24(1985). ACKNOWLEDGMENT S The authors thank Professor F. Horii and Mrs. K. Omine of Kyoto University for their helpful advice and assistance in the NMR measurements. Fmancial support by the Asahi Glass Foundation is gratefully acknowledged.


TRANSFORMATIO N OF BIOACTIV E GLAS S GRANULE S INT O CA-P SHELL S IN VITR O S. Radin, P. Ducheyne, S. Falaize, A. Hammond Department of Bioengineering, University of Pennsylvania 3320 Smith Walk, Philadelphia, PA 19104

ABSTRAC T Bioactive glass (BG) reactions were modeled in physiological solutions in static or dynamic conditions using either integral (no solution exchange) or differential (with solution exchange at designated time periods) modes of immersion. BG granules (either 300-355 or 200-300 |im) were immersed in tris buffered solution complemented with plasma electrolytes (TE), or with both plasma electrolytes and 10% serum (TES-10). Post-immersion solutions were analyzed for changes in Siconcentrations. Granules were analyzed for compositional, morphological and structural changes resulting from immersion by using scanning electron microscopy (SEM), energy dispersive X-ray (EDX) analysis and Fourier Transform Infrared (FTIR) spectroscopy. The amount of Si released from BG granules in the differential mode of immersion in TE was double the one in the integral mode. The dissolution was further enhanced in serum-containing TES-10: total dissolution of silica was observed after 1 week of differential immersion in TES-10. The granules with totally dissolved silica were transformed into hollow Ca-P shells. KEYWORDS : bioactive glass, in vitro, surface modification, dissolution INTRODUCTIO N A unique phenomenon of excavation of bioactive glass granules of narrow size range (300-355 |im), concomittant differentiation of osteoprogenitor cells, and formation of new bone tissue inside the excavated granules was observed in vivo [1]. Notwithstanding the biological evidence for this resorption to be cell mediated [1], solution mediated effects can also play a role. Before, bioactive glass (BG) reactions were modeled in physiological solutions using static or integral conditions (no fluid flow or solution exchange) [2,3]. However, physiological fluids in the body are in constant circulation. Thus, in this study, we modeled the reactions of BG granules in dynamic conditions using a differential mode of immersion. This immersion methodology represents conditions in which the solution is continuously replenished. In this paper we document that only differentially performed immersion experiments simulates the internal resorption process. MATERIAL S AND METHOD S BG 45S5 granules with a composition 45% Si02, 24.5% Na20, 24,5% CaO and 6% P2O5 (Orthovita, Malvern, PA), were immersed, either integrally or differentially, in tris buffered solution complemented with either plasma electrolytes 45

46


(TE) or complemented with both electrolytes and 10% newborn bovine serum (TES-10) at 370c for up to 1 week. Granules of either 300-355 or 200-300 |iim were used. The granules were immersed at a weight-to-solution ratio of 0.5 mg/ml. In the differential mode, the solutions were exchanged at 3, 6, 10, 24 hours and then every day. Post-immersion solutions were analyzed for changes in Siconcentrations using atomic absorption spectrophotometry (AAS, 5100, PerkinElmer, Norwalk, CT). Following immersion, the BG granules were analyzed for compositional and morphological changes using SEM/EDX (JEOL 6400) analyses. Morphological changes were viewed on granules attached to a carbon tape and coated with carbon. Some of the post-immersion granules were fractured under a light stereo microscope to expose the inner surface. Cross sectional analysis was performed on post-immersion granules embedded in epoxy and polished using diamond coated discs. After polishing the cross-sections were coated with carbon. FTIR analysis was also used to determine the structure and composition of the reaction surfaces. RESULT S AND DISCUSSIO N The Si-release from BG granules of 300-355 |Lim (expressed as weight % of the original Si content) vs. immersion time in the integral mode in TE, and the differential mode in TE and TES-10 is shown in Figures 1 a,b. Whereas 34.5% of the original Si-content was dissolved after 1 week of immersion in the integral mode in TE, the cumulative Si-dissolution in TE was equal to 69.6% in the differential mode. The dissolution was further enhanced by differential immersion in TES-10: total dissolution of the silica from granules was observed after 1 week of differential immersion. Similar results (i.e. total silica dissolution after differential immersion in TES-10) were obtained for BG granules of 200-300 ^m.

80- .

Le 40 -

1

20 ^

0 H f 3

-

- TE.inligr«

1^ 1 24

j

X

1 72

--

4

^

168 0

24

Time of immersion (hours )

Figure 1 a,b. Si-release (% of the original Si-content in BG) from BG granules (300-355 \xm)vs. immersion time in integral mode in TE (a) and differential mode in TE and TES-10 (b). Arrow indicates time to formation of a HA surface layer detectable by FTIR.

In Vitro Transformationof Bioactive Glass Granules (300-355 fiMj into Ca-P Shells: S. Radin 47

Figures 2 a, b. SEM micrographs of BG granules after differential immersion in TES-10 for 168 hours: (a) view of a fractured granule after silica dissolution; (b) cross-section of an internally dissolved granule (embedded in epoxy and polished). The SEM micrographs of granules after differential immersion in TES-10 (Fig. 2 a,b) show transformation of the granules into hollow shells. The EDX spectra (Fig. 3) corresponding to three different spots of the cross-section indicate that the hollow particle was mainly composed of a Ca-P phase which contained traces of CI, Na and Mg. The Ca/P ratio in the different spots on the shell varied from 1.1 to 1.3. No significant Si-content could be detected on either outer or inner surfaces of the shell. FTIR spectrum of the granules after differential immersion in TES-10 showed that the Ca-P phase was partly or fully crystalline, as indicated by the sharpening and splitting of the P-O band in the lower energy region (Fig. 4). The crystalline phase could not be identified because of extra absorption bands present in the energy region from 700 to 950 cm-1 due to spectral distorsion associated with a rough surface (i.e. small particles).

keV

Figure 3.Three EDX spectra (the spectra overlap) taken from three different spots on the hollow shell cross-section (shown on Fig. 2b): at the outer surface, at subsurface and on the inner surface.

48

Bioceramics Volume 10

Figure 4. FTIR spectrum of BG granules after differential immersion in TES-10 for 168 hours. Split of the P-0 band in the lower energy region indicates formation of a crystalline Ca-P phase. The observed effect of the mode of immersion and the presence of serum protems on the the degree of silica dissolution can be explained as follows. In the mtegral mode of immersion in serum-free TE silica dissolution was slowed down after 24 hours of immersion due to formation of a HA surface layer (detected by FTIR). The protective effect of the surface HA layer on dissolution of silica-based bioactive glass has been described elsewhere [4]. Our data indicate that the differential mode prompted continuous Si dissolution in TE at the immersion stages preceedmg the HA formation (Fig. la). As we previously reported [5], the presence of serum in TES-10 slows down the HA formation. The surface reaction layer which forms in TES-10 is composed of amorphous Ca-P accumulations in a silica matrix. This layer does not prevent continuous Si-dissolution. As a result all the sihca dissolved from the granules in these experiments. Granules of which all the sihca was dissolved in TES-10, were fully transformed to hollow Ca-P shells. Crystallization of the Ca-P phase occured upon Si-dissolution. CONCLUSION S 1. In vitro immersion in a serum solution using conditions reflecting the continuous fluid flow in vivo, leads to internal dissolution of glass granules. Static immersion experiments fail to achieve this phenomenon. 2. Transformation of bioactive glass granules into hollow Ca-P shells can be achieved in vitro. AKNOWLEDGMEN T This work was supported by grants NSF BSC 93-09053 and NIH DE-10693

REFERENCES

1. Schepers, E., De Clercq, M., Ducheyne, P. and Kempeneers, R., J.Oral. Rehab..,1991, 18, 439-452. 2. Hench, L.L., J. Am. Cer. Soc, 1991, 74(7), 1487-510. 3. Kokubo, T., J. Non-Cryst. Solids, 1990, 120, 138-151. 4. Hench, L.L., J. Non-Cryst. Solids, 1978, 28, 83-105 5. Radin, S., Ducheyne, P., Rothman, B., Conti, A., J. Biomed.Mat. Res., 1997, 35, in press.


Multilayere d Coating s of Hydroxyapatite/Glas s Cerami c Composite s Plasm a Spraye d on Ti-6A1-4V Alloy P.L.Silva*^* J.D.Santos^*, F.J.Monteiro^* ÎSEP-CEIA-Instituto Superior de Engenharia, Institute Politecnico do Porto, R. de S.Tome, 4200 Porto, Portugal. INEB-Instituto de Engenharia Biomedica, Praâ Coronel Pacheco 1,4050 Porto, Portugal. ’-’Departamento de Engenharia Metalurgica, FEUP, Universidade do Porto, R. dos Bragas, 4099 Porto Codex, Portugal. ABSTRACT Aiming at obtaining plasma sprayed coatings with enhanced bioactivity, when in contact with host tissues, double layered coatings were prepared. The top layer was composed of a hydroxyapatite (HA)/4% P2O5 glass composite and the undercoat was of pure simple HA onto Ti-6A1-4V alloy. X-Ray Diffraction Analysis (XRD) of the composite top layer showed that it was less crystalline than simple HA layer due to the glass addition, and was composed of a HA matrix with a small amount of p-TCP. Scanning Electronic Microscope (SEM) observations showed that it has been possible to obtain double-layered coatings with good bonding between them and well adherent to the metallic substrate. Coating to substrate adhesion was determined according to three standard methods: ASTM C 633-79, ASTM D 1002 and DIN 50161. Image analysis technique was used to determine the type of failure mechanism that took place and it was found that 85.6%–2.7 was adhesive and 13.4%–2.7 cohesive. The coating to substrate adhesion values are within the range commonly required for this type of plasma sprayed coatings for medical applications. KEYWORDS : Plasma-spraying, ceramic composites, adhesion. INTRODUCTIO N The interest in using HA for biomedical applications is commonly known [1-3]. However, it has been proved that for load bearing applications the mechanical properties of HA, mainly its strength and fracture toughness should be improved [4-5]. In a previous work [6], it has been shown that by applying a P2O5 based glass reinforcement to HA, fracture toughness and biaxial bending strength were enhanced. In this work, HA reinforced with 4%P205 based glass was used as a surface layer, i.e., the material that primarily establishes contact with tissues and organic fluids. In order to adjust glass chemical composition as closely as possible to inorganic part of bone chemical composition [7-9], Na20, MgO and K2O oxides were incorporated into P2O5 based glass. This composite was the top layer of a mixed coating containing as an undercoat a pure HA layer, adherent to the Ti-6Al-4V alloy. One of the main concerns when using plasma sprayed coatings is to obtain adequate coating to substrate adhesion [10]. The way by which a coating adheres to a substrate is very complex and not fully understood [11-13]. Several theories and mechanisms of adhesion have been proposed but there is no single one that might explain all adhesion behavior [11-14]. 49

50


P205

35

Table 1 - Glass chemical composition (mol%). CaO Na20 K2O 35 10 10

MgO 10

Due to the difficulty in quantifying this property, three standard methods have been chosen allowing for crossed results evaluation. MATERIAL S AND METHOD S P205-CaO based glass was prepared from reagent grade chemicals and its chemical composition may be seen in Table 1. The glass was wet mixed and milled with HA powder using methanol. 4% glass addition was used. Mixed powders were dried and isostatically pressed at 200MPa, The method used for composite preparation has been fully described elsewhere [6]. Cylindrical shaped samples were sintered at 1300^C and once again milled, using an agate ball mill pot. The powders were then sieved to obtain a grain size distribution suitable for plasma spraying deposition. Two kinds of HA were used for the base coatings: HA-P, supplied by Plasma-Technik with the reference code, AMDRY AM 6021, specifically prepared for plasma spraying and used "as received"; HA-I, supplied by Plasma Biotal and conditioned to be used for plasma spraying. Plasma spraying was performed under atmospheric conditions using a Plasma Technik automated equipment. After grit blasting the Ti-6A1-4V alloy with AI2O3 spheres, and chemical degreasing with trichlorethylene, a 50\im HA coating was sprayed followed by the deposition of SOjxm HA/glass composite coating. Samples were characterized by XRD and compared with an XRD spectrum obtained for a 100|Lim commercial HA-P coating. Ten samples were used for each of the three adhesion experiments. From ASTM D 1002 tests, comparative shear strength values were obtained. With the ASTM C 633-79 it was possible to obtain coating adhesion under stresses normal to the surface, and with ASTM C 633-79 DIN 50161

0 1,5+0,2 cm

l . t cm

ASTM D1002

1,6mn n 1

1 Z.TTnm

sttear

"^.âfcr^ ^

Fig.l - Test specimens according to ASTM D 1002, ASTM C 633-79 and DIN 50160 standards.

Coatings of HA/Glass Ceramic CompositesPlasma Sprayed on Ti-6al-4v Alloy: P.L. Silva et al.

v._^LÛ--^X^w ;

51

M- MU AMkli&«ii^.^.v^^^^ ^

a) b) Fig.2-XRD spectra for HA-I+composite samples a) and HA plasma sprayed samples b). DIN 1061 testing compressive stresses were applied. The shape and dimensions of the test specimens may be seen in Fig.l. The glue used to assemble the test specimens in both, ASTM D 1002 and ASTM C 633-79 tests, was Plasmatex Klebbi from Plasma Technik. In order to determine what type of failure mechanism took place, cohesive or adhesive, fracture surfaces were observed by SEM according to image analysis technique. RESULT S AND DISCUSSIO N XRD spectra may be observed in Fig.2. Both coatings had an HA matrix with small amounts of P-TCP. Comparing with the HA + composite samples, the simple commercial HA coating presents higher cristallinity. Adhesion results are presented in Table 2. Slightly higher adhesion values were obtained for the HA-I+composite samples. Both coatings have shown higher resistance to uniaxial tensile stress than to shear stresses, applied in ASTM D 1002 experiments or compressive stresses used in the DIN 60161 standard. Examples of image analysis applied to the samples tested according to ASTM C 633-79 are presented in Fig.3. The areas where some coating remained attached to the substrate are yellow colored and the gray color represents Ti-6A1-4V alloy substrate. Using this technique it was possible to determine that HA-I+composite failure was 85.6 – 2.7% adhesive and 13.4% – 2.7 cohesive. HA-P+composite samples have shown 57.2%–4.6 adhesive and 42.8%–4.6 cohesive failure. In samples subjected to ASTM D 1002 the coating was completely transferred to the counterpart so that image analysis determinations were not performed. Fracture surfaces of samples subjected to DIN 50161 have showed that failure occurred by the coating - cohesive failure. Magnification of the failure area have shown that fracture surfaces were irregular and some failure planes may be observed, which indicates that brittle failure took place. SUMMAR Y In this work an attempt was made to obtain new coatings designed to tailor the needs of bonding to neighbor tissues and whitstanding for a reasonable amount of time the contact with physiological environment. A surface coating capable of inducing a fast response from the host tissue in the early moments of implantation was created, showing a more reduced dissolution kinetics in the long term, thus allowing for the newly formed bone tissue to be fully established Coating HA-I+composite HA-P+composite 1 Glue

ASTM D 1002 (MPa) 14.9–4.3 13.1–3.8 22.7–0.1

ASTM C 633-79 (MPa) 35.4–6.5 33.8–5.6 70.0–0.1

DIN 50161 (MPa) 40.0–9.8 36.1 –8.6 2

1

52


Fig.3 - Failure surface image analysis photomicrograph of an HA-I+composite sample a) and HAP+composite sample b) and adequately fixed to the implants. As it was clearly detected in Fig.2, the upper layer coating was more amorphous than HA, probably due to the glass incorporation. Results presented in Table 2 show that coating to substrate adhesion was slightly improved when HA-I + composite double layers were used. These results may be explained by the fact that HA powders used to produce these coatings had higher strength and density than those used to prepare the HA-P+composite coatings since they were isostatically pressed and sintered. Both kinds of double layered coatings present higher adhesion values for ASTM C 633-79 - tensile stresses than for ASTM D 1002- shear stresses or DIN 50161 - compressive stresses. Comparing these results with the needs for the implanted materials in load bearing applications, 30MPa, it may be concluded that both double layers seem to entirely satisfy stress requirements. Image analysis, seen in Fig.3 and performed on samples subjected to ASTM C 633-79 tests, have shown that adhesion failure was 85% adhesive for HA-I + composite samples but, for HAP+composite coatings this value decreases to 57%. This behavior may be explained by the fact that when using HA-I+composite samples both HA - the one from the composite layer and the "pure" HA layer - were prepared by the same procedure. HA-P+composite samples were prepared with two different kinds of HA powder, the HA powder from the composite that was isostatically pressed and sintered, and the commercial HA from the bottom layer that was applied "as received". The fact that in the ASTM D1002, the double layer was completely transferred to the counterpart have showed that this was a mainly adhesive failure. The DIN 50161 failure surface observations presented essentially cohesive kind of failure.

Acknowledgment s The authors which to thank INEB for the provision of laboratory facilities, financial support of JNICT trough Ref.PBICT/CTM/1890/95 and Marta Sa and Barbara Silva for their immense collaboration in the execution of the adhesion experiments.

REFERENCES 1. 2. 3. 4. 5. 6. 7. 8. 9. 10. 11. 12. 13. 14.

RDucheyne , J.BiomedMater.Res.App.Biomater., 21 (1987), 219. H.Aoki, in ScienceandMedical Applications of Hydroxy apatite, Takayam a Pres s System Center , JAAS, Tokyo, (1991), 165. R.Z.Legeros , Adv.Dent.Res.,2, (1988), 164. G.With,H.Va n Dijk, N.Hattu , K.Prijs , J. Mater.Sci., 16, (1981), 1592. M.Akao , H.Aoki, K. Kato , J. Mater.Sci., 16, (1981), 809. J.D. Santos , P.L. Silva, J.C . Knowles, F.J.Monteiro , /. Mater.Sci., Mater.Med.,1, (1996), 187. D. Williams , Medical& Dental Materials, Pergamo n Press , (1986). C.Rey, Biomaterials, 11,(1990), 14. G.Evans , J.Behiri , J.Currey , B.Bonfield, /. Mater.Sci., Mater.Med.,1, (1990), 38. D.Matejka , B.Benko, Plasma Spraying ofMetallic Materials, Joh n Wiley & Sons, (1989). S.D.Brown , ThinSolidFilms,119, (1984), 127. K.L.Mittal , Adhesion Me^LSurement of ThinSolidFilms,Thick Films and Bulk Coatings , ASTM STP 840. R. Dixon, Surface Engineering, 11,(1993), 4. W.Lian , Novel Plasm a Sprayin g Processin g for Enhance d Surfac e Engineering , Matetrial s World , 4,


THE BONY REACTIO N TO RAPIDL Y DEGRADABL E GLASS-CERAMIC S BASED ON THE NEW PHAS E Ca2KNa(P04) 2 C. Muller-Mail, G. Berger^, C. Voigt^ B. Bakki^, U. Gross^ Department of Traumatology and Reconstructive Surgery^ and Institute of Pathology^, Universitatsklinikum Benjamin Franklin, Freie Universitat Berlin, Hindenburgdamm 30, 12200 Berlin, Germany, Federal Institute for Materials Research and Testing, Berlin-^, Germany. ABSTRAC T Glass-ceramics in the composition field of CaO-K20-MgO-Na20-P205 can be produced based on the main crystalline phase Ca2KNa(P04)2. The degradation rate depends on the whole composition, i.e. the crystalline phases as well as the residual glass-phase. The degradation rate of the two investigated glass-ceramics was higher as in B-tricalciumphosphate (B-TCP) and was mainly due to passive leaching processes, especially in GB18. Both materials exhibited bonebonding properties and allowed at least for partial guided bone regeneration during degradation. KEYWORDS : Bone/interface/glass-ceramics/degradation/guided bone regeneration/ultrastructure. INTRODUCTIO N Bone defects, for example after trauma, need reconstruction. Such defects can be reconstructed by using autologous bone and homologous bone, both having their drawbacks and limitations. An alternative method uses bioactive bone-bonding implant materials which display different degradation rates. In most of these materials the interface is stabilized by bone-bonding and therefore, the bone bonding material is not further subject to degradation and replacement by bone [1]. On the other hand, the higher degradable materials should allow for guided bone regeneration while being degraded. Thus, the degradation rate should not be too high. The goal of this study was to evaluate two new synthetic implant materials based on a Ca2KNa(P04)2-phase as a glass/glassceramic composite. The dissolution rate of these materials in comparison to 13-TCP is enhanced by exchanging Ca against other ions such as K, Na and by adding Mg [2]. MATERIAL S AND METHOD S Two particulate implant materials were tested with the following constituents according to [2]: GB14 with a content of CaO 30.67 %, MgO 2.45 %, Na20 9.42 %, K2O 14.32 % and P2O5 43.14 % (particle dimentions were measured to be 370 – 70 |nm width and 610 – 100 jam length, n = 45), and GB18 containing CaO 17.72 % and MgO 12.75 %, Na20 9.79 %, K2O 14,88 %, P2O5 44.86 % (380 – 70 |im width, 630 – 100 |Lim length, n = 27). Both compositions are meltable and crystallize spontaneously fi-om the melt. Thus, they can be easily produced. Both particulates consist of a main crystal phase and other amorphous and crystal moieties. The dissolution rate was measured in vitro (0.2 M TRIS-HCl) and yielded rates of 292 – 40 and 1278 – 80 mg/1 for GB14 and GB18, respectively. Therefore, the materials were much more soluble as B-TCP in the same testing system (30,6 – 10 mg/1)[3]. The dissolution rate is controllable by the amount of added ions, such as Na, K, or Mg. Particulates were sterilized by dry heat at 180 C for 30 min.. 100 mg 53

54


were implanted into each femur below the patella sliding plane into trabecular bone in Chinchilla rabbits after arthrotomy of the knee joint. The defect containing the particulates was closed using a chondrocortical bone slice produced by the hollow cylinder drill after removing the adherent trabecular bone below the cartilage. After implantation, this slice was fixed in the anatomical position by the pressure of the revised patella. Thus, no particles were observed in the knee-joint. A total of 9 animals was operated bilaterally per material and each 3 animals were sacrificed at 7, 28, and 84 days and the femora were prepared for scanning (SEM) and transmission (TEM) electron microscopy as described [4] except that the specimens for SEM-investigation were air dried after rinsing in hexamethyldisilasane. RESULT S AND DISCUSSIO N Specimens prior to implantation In the SEM the GB14-particles showed three different appearences. Two particle types consistend mainly of crystalline spheres attached to each other by a thin film of amorphous matrix. The crystalline spheres had diameters of either 10 – 1 |im (n = 30 measured) or 45 – 6 )Lim (n = 30). Between the crystalline spheres representing the ceramic moiety and the amorphous glass-phase there were some cracks up to 2 |Lim width and more than 20 |im length. At the neck of separate spheres, where cracks contacted each other, roundish pores were created of approximately 3.5 |Lim diameter. The third kind of particles was the most dense, since there were no cracks or pores and crystalline spheres were not detectable. The GB18 particles were also dense and homogeneous with only few deepenings and elevations. There were no apparent substructures (Figure 1). Specimens after implantation All cases showed healing without special events. During the time course between 7 and 84 days of implantation the surface morphology of the three GB 14 particle types changed considerably. The crystalline spheres either of 45 or 10 jam in diameter were degraded, leading in the case of 45 )Lim spheres to a microporosity of their surfaces with pores and cracks with diameters of 1 jiim or 2 \xm width, respectively. Later on, this process led to a destruction and formation of smaller substructures (Figure 2). Thus, irregular porous substructures of some |Lim in length were produced indicating a rapid degradation of the spheres. The 10 \im crystalline spheres decreased contineously in size by keeping their spherical outline. In the third kind of GB14 particles the amorphous moiety was leached at already 7 days liberating roundish crystalline substructures with mean diameters of 13 – 4 |Lim (n = 10). Between the amorphous moiety and the spheres cracks developed which increased in width at longer time intervals (Figure 3). Already at 7 days pores of about 0.5 fim developed in the spheres. The pore diameters grew with time leading to a complete destruction in some cases. In vitro experiments have shown, that this degradation process led to high concentrations of Mg and P in the medium as well as to precipitations rich in these ions on the substrate. The osteoblast growth was slightly inhibited as compared to Thermanox^[5]. In the GB18 particles at 7 days the ceramic moiety consisting of rod-like structures of up to 20 fim length was liberated, possibly by leaching of the surrounding amorphous glass moiety. Later on, these rod-like structures were also degraded as in the GB14 particles and smaller irregular substructures were formed (Figure 2). At 7 days there was almost exclusively fibrous material surrounding the different particle types. At 28 days bone was observed in contact to the GB14 particles. The bone contacted the dense particles and the particles with 45 \xm crystalline spheres directly, whereas the particles with small 10 \xmspheres were separated from the bone by clefts of about 1 |um width. The different tissue

Bony Reaction to Degradable Glass-Ceramics Based on the Phase Ca2KNa(P04)2: C. Muller-Mai et al.

55

reactions indicate that the chemical composition of the different moieties within single particles as well as in particle types was varying. In general, GB18 developed more bone contact as GB14, probably related to the higher reactivity. There was bone at 28 days in contact to the particles and at 84 days there was almost complete bone contact of the GB18 particles. In the TEM, all particulates exhibited crystalline and amorphous parts. There seemed to be empty spaces between both phases suggesting some microporosity prior to implantation. At each time interval there was soft tissue with fibroblasts and macrophages as well as bone in the interface with increasing bone amounts at longer time intervals especially in contact to GB18 and the dense GB14 (Figure 3). Mineralization occured in the cracks between spheres. Both particulates were degraded with a higher rate of the GB18. In GB14 this was due to a destruction of the spheres as well as to leaching of the amorphous moiety. In GB18 there was first leaching of the amorphous phase leading to the liberation of rod-like structures which degraded also later on. REFERENCE S 1. Muller-Mai, CM., Voigt, C , Baier, R.E. and Gross, U. Cells & Mater. 1992, 4, 309-327. 2. Berger, G., Gildenhaar, R. and Ploska, U. In: Bioceramics Volume 8, Wilson, J., Hench, L.L., Greenspan, D. (eds), Pergamon, Oxford, UK, 1995, 453-456. 3. Berger, G., Gildenhaar, R., Gross, U., Knabe, C, Loginow-Spitzer, A., Miiller-Mai, C , Ploska, U. and Radlanski, R. In: Werkstoffe fiir die Medizintechnik, Symposium 4, Breme, J. (ed), DGM-Informationsgesellschaft Verlag, Frankfurt, Germany, 1996, 59-65. 4. Miiller-Mai, C , Voigt, C. and Gross, U. Scanning Microsc. 1990, 4, 613-624. 5. Knabe, C , Gildenhaar, R., Berger, G., Ostapowicz, W., Fitzner, R., Radlanski, R., Gross, U. and Siebert, G.K. In: Transactions, Vol. I, 5th World Biomat. Congress, International Liaison Committee (ed), University of Toronto Press, Toronto, Canada, 1990, 890.

Figure 1: Surfaces of the particle types prior to implantation. A: GB14 with 45 |Lim spheres, cracks and a pore; B: GB14 with 10 |Lim spheres; C: Dense GB14; D: GB18. SEM, bar 20 jim.

56


Figure 2: A: Spheres (45 jim-type) at 7 days with newly developed pores and cracks. B: GB18 at 84 days surrounded by bone with rod-like crystalline moiety. SEM, bars A:50 |Lim, B: 100 |Lim.

Figure 3: A: Dense GB14 at 84 days with cracks between amorphous and crystalline moiety with pores in separate spheres surrounded by bone. B: Bone bonding to ceramic moiety of GB18 at 84 days (grey, white areas as artefacts due to loss of glass-ceramic). TEM, bars A: 100 |im, B: 1 jim.


RESORBABL E POROU S PHOSPHAT E INVER T GLASSE S - FIRS T IN VITR O AND IN VIVO RESULT S J. Voger, K.-J. Schulze^ D. Reif, P. Hartmann^ U. Platzbecker^ B. Leuner^ ’Otto-Schott-Institut, Friedrich-Schiller-Universitat, FraunhoferstraBe 6, 07743 Jena, Germany, ^Klinik fiir Orthopadie, Technische Universitat, FetscherstraBe74, 01370 Dresden Germany, ^Biovision GmbH, Am Vogelherd 1, 98693 Ilmenau, Germany, ’’Institut fiir Optik und Quantenelektronik, Friedrich-Schiller-Universitat, Max-Wien-Platz 1, 07743 Jena, Germany

ABSTRAC T A resorbable phosphate invert glass was investigated in view of its solubility behavior "in vitro". First animal experiments were carried out using a porous shape of the phosphate invert glass and of the corresponding glass ceramic. Both the solubility tests and the animal experiments confirm an excellent biocompatibility. The living bone grows into the pores of the implants and the materials are reabsorbed in course of time. The resorption rates meet the expectations. KEYWORD S Phosphate glasses, resorbable, porous, solubility tests, animal experiments INTRODUCTIO N The regeneration of bony defects can be supported by resorbable implants. Sintered tricalcium phosphate is mainly used for these clinical applications today. Pure dense sintered TCP shows a good biocompatibility [1,2]. However, in dependence on size and place of the filled bony defect, the resorption term of this material can reach up to some years and rests of the material can cause biomechanical problems. Additionally, the different solubilitys of the crystalline phase and the amorphous grain bounderys can result in a disintegration of the material in grainy particles. In these cases, a biocompatible resorbable homogeneous material possessing a clearly increased resorption rate is desirable. Some calcium phosphate invert glasses are suitable for the development of resorba› ble implants showing well defined resorption rates [3,4]. Because these glasses don’t show phase separation, they represent single phase systems with an uniform solubility. Using a salt-sintering process, the glasses can also be produced in a porous shape with a texture similar to that of spongious bone. A selected glass of this group and the corresponding glass ceramic were used for animal experiments and solubility tests. 57

58


MATERIAL S AND METHOD S The composition of the investigated glass/glass ceramic was determined by usual wet chemical methods: 32.6 mol% P2O5, 27.6 mol% CaO, 27.6 mol% Na2 0, 12.2 mol% MgO. Mixed sodi› um/calcium diphosphates and a B-TCP phase, stabilized by magnesium, are the main crystal phases of the glass ceramic. ^^P MAS NMR experiments confirmed the invert glass structure of the materials. They contain not or low condensed phosphate groups exclusively. In their porous form both materials show an total open porosity of 65%. The diameters of the pores vary between 150|Lim and 400^m (figure 1). Smaller pores in the walls between the main pores are caused by the sintering process and range from about l|im and 60^m (15%). Informations about the solubility behavior in dependence on time were obtained using a simple soaking test: Two grams of the substances (grain size 315|Lim to 400 |im) in 200 ml destilled water were shaked by a shaking machine at 37 C for 24 hours. In the eluate, the pH-value and the quantity of dissolved ions were determined. The soaked glass grains were washed and the same procedure was repeated eight times. For the animal experiments guinea pigs were used. Cubes (2x2x2 mm^) of the porous materials were implanted into the tibiae of 72 animals, both sides. After 2, 4, 8, 16, 32 and 64 weeks bone speci› mens containing the rests of the implants were resected. Besides the histological investigation, also histomorphometrical methodes were used. The determination of the total area of the implants by a light microscope combined with a drawing unit and a PC gave informations about the resorption rate in vivo. RESULT S AND DISCUSSIO N Table 1 gives the total quantity of dissolved ions (E Ca^^Mg^^,Na^,P04 în mg/1) in dependence of soaking period. The values of the glass are five to seven times higher than the values of TCP. Because of the strongly increased surface of the porous glass also the quantity of dissolved ions is

Figure 1. Resorbable porous phosphate invert glass with a total open porosity of 65%.

Resorbable,Porous PhosphateInvert Glasses: J. Vogel et al.

59

Table 1. Total quantity of dissolved ions (E Ca^^,Mg^^,Na’^,P04 ^’in mg/1) in dependence on soaking period eluatel

eluate2

eluateS

eluate4

eluateS

eluate6

eluate7

eluate8

glass

115.5

126.3

117.3

85.6

89.1

84.8

83.8

78.5

porous glass

290.1

190.1

144.8

120.0

112.6

153.5

160.7

159.0

TCP

16.0

17.8

13.8

15.8

16.5

increased. The reason for the minimum at the fourth and fifth period is unknown in this time. Possibly, an unstable intermediate layer is formed because of the higher reactivity of the porous material. The delayed resorption in vivo some weeks after implantation (see below) could be caused by such a layer. However, the amounts of the several dissolved ions never exceed the physiological tolerance. After an initial period, the values of both glass and porous glass are nearly constant. This is also valide for the pH-values of the eluates (figure 2). Similar to TCP the pH-values come close to the physiological pH-value. Therefore, cell toxicological problems are not to be expected. The lower pH-values of the porous glass in the first eluates is caused by the production process of the porous texture. The histological investigations of both glass and glass ceramic didn’t show any symptoms of inflammation. Within the first two weeks post operation the formation of thick osteoid beams was observed. In the following time the pores act as a guide rail for the young bony cells growing in and the implants are completely incorporated by osteoid after three to four month.

10,Q|

physiologica l pH TCP porous phosphat e glas s phosphat e glas s

9,59.0 i 8.5-

X

Q.

8,07,5 ^

7.oH

,,>^J ;

6,56,0 -

4

5

6

eluate number

Figure 2. pH-values of eluates in dependence on soaking period

60


10Oi 4

V

\

^

80-

^ ^ ?

60-

^

40-

A

0\ \ ’^^

20-

()

glass cerami c phosphat e glass j

A>.....___ ^

1

10

20

30

1

40

1

50

1

60

weeks

Figure 3. Total area of resorbable implants in vivo in dependence on time From that time on the osteoblasts of the osteoid are changed to osteocytes. After 64 weeks the spongiosa framework represents a mixture of ripe bone, osteoid and small amounts of incorporated glass/glass ceramic particles. Simultaneous to the formation of bone the porous phosphate glass and glass ceramic are reabsorbed. Figure 3 gives the course of the total area of the implants in dependen› ce on time. 64 weeks post operation only small rests of the materials are detectable (glass: around 2%, glass ceramic: around 10 %). A mechanical instability caused by the dissolution of the implants was not observed. Obvoisly, it is compensated by the increasing bony integration of the implants. CONCLUSION S Both the phosphate invert glass and the glass ceramic meet the requirements of resorbable implant materials. Simultaneous to the resorption process the implants are bony integrated. The resorption rate of the materials is adjusted to the growth and the mineralization of bone. REFERENCE S 1.

Koster, K., Heide, H., Konig, R., Langenbecks Arch. Chir. 343 (1977), 174

2.

Klein, C , Driessen, A., de Groot, K., van den Hooff, A., J. Biomed. Mater. Res. 17(1983), 769784

3.

Hartmann, P., Vogel, J., Schnabel, B., J. Non-Cryst. Solids, 176 (1994), 157-163

4.

Vogel, J., Hartmann, P., Schulze, K.-J. In: Advances in Science and Technology 12: Materials in Clinical Applications, Techna Sri., Faenza 1995, 59-66


IMPLANTATIO N OF BIOACTIV E AND INER T GLAS S FIBRE S IN RAT S TISSU E RESPONS E AND SHORT-TER M REACTION S OF THE GLAS S

SOFT

M. Brink\ P. Laine^ K. Narva^ and A. Yli-Urpo^ ^Department of Chemical Engineering, Abo Akademi University, Biskopsgatan 8, FIN20500 Abo/Turku, Finland Înstitute of Dentistry, University of Turku, Abo/Turku, Finland ABSTRAC T The purpose of this work was to develop a bioactive glass fibre that resorbs in soft tissue without causing inflammatory reactions. In addition, the glass should bond to bone and be easily manufactured. Two different biocompatible glasses were chosen for implantation, and glass surface reactions as well as soft tissue response were evaluated. An inert commercial glass fibre was used as reference. After implantation, all glasses were in good contact with the surrounding tissue. The biocompatible glasses were severely resorbed after 28 days in soft tissue indicating that these glasses are suitable for membranes in orthopaedic and maxillofacial surgery, and for reinforcement of resorbable biopolymers. The reference glass fibre did not show any signs of reaction. KEYWORD S glass fibre, glass reactions, resorption, soft tissue, tissue response INTRODUCTIO N Resorbable bioactive glass fibres may be used as membranes for tissue guiding and as carriers for growth factors in orthopaedic and maxillofacial surgery. The glass fibres may also be used for reinforcement of resorbable biopolymers. For these applications, the fibres should preferably resorb within weeks after implantation without causing any inflammatory reactions. In addition, the fibres must be easily manufactured without risk for devitrification (crystallisation) of the glass melt or fibres. Pazzaglia et al. [1] have developed bioactive glass fibres for substrates for bone apposition, but implanted into rat soft tissue, an intensive inflammation reaction occured. Other bioactive glass fibres, but intended for bone implants, have been presented by Vita Finzi Zalman et al. [2] while Graves and Kumar [3] have developed a bioabsorbable glass fibre for reinforcement of bioabsorbable polymers. The latter fibres are based on the system CaO-P205.

Two different glasses, glass 20-92 and 13-93, were selected for implantation into rat soft tissue. Glass 20-92 is biocompatible but since it contains only 50 wt % SiOi, it has a narrow working range [4,5]. Glass 13-93, containing 53 wt % Si02, is bioactive and has a large working range [4,5]. An inert commercial glass fibre, the E-glass, was used as reference material. After implantation in rat soft tissue, initial glass reactions as well as soft tissue response were evaluated. 61

62


MATERIAL S AN D METHOD S The denotations and compositions of the fibres used are given in Table 1. The glasses 20-92 and 13-93 were melted for 1.5 h at 1360’’C in a ceramic crucible (Hackman-Arabia, Finland). The batch size was about 300 g, and all raw materials except the sand were of analytical grade. Glass fibres were obtained by pouring the glass melt onto a rotating plate (about 30 cm in diameter) of stainless steel. The thickness of the fibres could be controlled by changing the spinning velocity. Fibres of glass 20-92 are shown in Figure 1. The E-glass fibres were obtained from Ahlstrom Glass Fibres, Karhula, Finland. The sizing was removed by heating for 40 minutes at 650^C. Prior to implantation, all glass fibres were cut into a length of 5-7 mm, rinsed with ethanol and sterilised in hot air. Two or three different glass fibres were implanted subcutaneously into the soft tissue of 18 Long Evans male rats weighting 290-440 g. The rats were anaesthetised with 0.6-1.0 ml Hypnormfi/Dormicumfi. After the implantation times of 7, 14 and 28 days, the rats were killed with CO2. Tissue samples were fixed in 70% alcohol and embedded into plastics. Histological sections were made using a cutting-grinding technique and stained with toluidine blue. Tissue reactions were analysed with light microscopy and glass surface reactions by scanning electron microscopy (SEM) and energy dispersive X-ray analysis (EDXA). RESULT S AN D DISCUSSION After 7 days, no inflammatory cells were detected around glass 13-93 and 20-92. In general, the number of inflammatory cells around the glass fibres was minimal after all implantation times, and most of the surrounding tissue was inflammation-free. The number of inflammatory cells around E-glass was minimal after 7 days. All glass fibres were in good contact with surrounding tissue, and connective tissue grew in tight contact with the glass surfaces. Figure 2 and 3 present glass 20-92 and 13-93, respectively, after 7 days in rat soft tissue. For fibres of glass 20-92 and 13-93, it was found that the resorpfion had started already after 7 days in soft tissue. The resorption was detected as a silica rich layer at the fibre surface, with sporadic formation of calcium phosphate on top. This result is in accordance with previous studies on rods of glass 13-93 in soft dssue [6]. For glass 20-92, with a durability significantly lower than that of glass 13-93 [7], only a core of original glass surrounded by a silica layer was left after 7 days in vivo. The resorption of this glass was thus more pronounced than that for glass 13-93. However, the surrounding tissue did not show any signs of inflammation. The E-glass showed neither any inflammatory reaction in soft tissue, nor did it resorb. E-glass after 14 days in rat soft tissue is shown in Figure 4. After 28 days in rat soft dssue, fibres of glasses 20-92 and 13-93 were severely resorbed. Glass surface reactions after implantation are presented in Table 2.

Table 1. Denotations and composifions (in wt %) of the investigated glass fibres, and fibre diameter. Fibre 0 Glass NaiO K2O MgO CaO B2O3 AI2O3 P2O5 Si02 70-300 ^m 20-92 15 15 2 15 3 50 70-300 \im 13-93 6 12 5 20 4 53 10 |Lim** E-glass 1* 0.7 22.5 6.4 14.7 54 *Na20+K20 **The E-glass was implanted as a bunch of fibres.

Implantationof Bioactive and Inert Glass Fibres in Rats: M. Brink et al.

63

Table 2. Glass surface reactions after implantation into rat soft tissue for several observation times. (Ca,P = calcium phosphate) Glass 7 days 14 days 28 days 20-92 Silica gel with Ca,P Silica gel with Ca,P Silica gel with Ca,P 13-93 Silica gel with Ca,P Silica gel with Ca,P Silica gel with Ca,P No reaction No reaction No reaction E-glass* *without sizing CONCLUSION S This study indicates that glasses 20-92 and 13-93 are suitable for biomedical use as resorbable materials. Which of these two is the more suitable depends on the application. It was shown that fibres of these two glasses resorbed while the E-glass fibres did not. The resorption was initiated almost immediately but no inflammation reaction was detected. ACKNOWLEDGEMENT S This work was financially supported by the Finnish Technology Development Centre (TEKES) and the Academy of Finland (FA).

K-i^

Figure 1. Fibres of glass 20-92.

Figure 2. Glass 20-92 after 7 days in rat soft tissue, (magnification 125x)

64


Figure 3. Glass 13-93 after 7 days in rat soft tissue, (magnification 125x)

Figure 4. E-glass after 14 days in rat soft tissue, (magnification 125x)

REFERENCES 1. Pazzaglia U.E., Gabbi, C , Locardi, B., Di Nucci, A., Zatti, G. and Cherubino, P. J Biomed.Mater.Res. 1989, 23, 1289-1297. 2. Vita Finzi Zalman, E., Locardi, B., Gabbi, C. and Tranquilli Leali, P. WO 91/12032 i. Graves (Jr.), G.A. and Kumar, B. United States Patent 4,604,097. 4. Brink, M. /. Biomed.Mater.Res. (to appear). ^’ appe^)^" ^"’’"’^"’ ^ ’ "’’PP’"’^’’’ ^-^- ^ ^ Yli-Urpo, A. J. Biomed.Mater.Res. (to

’

TTL^r.’Ît

’

" ’ ^’’’""’^"^ ^- ’ " ’’^’’

^’’’’

BiomaterialsCongress,

7. Brink, M., Karlsson, K.H. and Yli-Urpo, A. WO 96/21628 (pending).

Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedingsof the10th InternationalSymposiumon Ceramics in Medicine, Paris, France, October1997) '1997 Elsevier Science Ltd

Spinal Fusion using Titanium Spacers with Bioglassfi and Autogenous Bone: A Comparative study in Sheep June Wilson\ Gary Lowery^, Stephen Courtney^ 1. Imperial College, Dept. of Materials, Prince Consort Rd. London SW7 2BP Tel/Fax 44(0)171 594 6745 2. Research Institute International, 6400 W. Newberry Road, Ste. 206, Gainesville, Florida 32605-4391 USA

ABSTRAC T Spinal fusion, or the surgical accomplishment of a bony union of deficient vertebral segments, is required in cases where long-term immobilization of these segments is indicated. In the anterior lumbar spine, this procedure requires the use of an intervertebral graft construct which is able to support the axial forces until confluent healing of the graft material occurs. In this study titanium surgical mesh spacers were filled with autologous bone, Bioglassfi, or a mixture of the two in the adult sheep spine. After three and six months the new bone was assessed both qualitatively and quantitatively and results show that although at three months autologous bone is to be preferred, by six months there is no significant difference between the three graft materials. This is particularly important since provision of sufficient autologous bone is always difficult and is associated with a high morbidity rate.

INTRODUCTIO N Spinal fusion is indicated in a variety of disorders of the lumbar spine. The goal of this surgery is to achieve bony immobilization of the affected levels. Anterior interbody fusions, usually relying on the use of autologous bone, are indicated in many cases of discal deficiency [1]. AAV ceramic monoliths [2] and titanium mesh spacers packed with autologous cancellous bone [3] have been used in this procedure. Availability and morbidity issues limit the use of harvested autologous bone [4], while the use of allograft involves inherent concerns regarding disease transmission. The ideal bone graft supplement to autologous bone should be both osteoinductive and osteoconductive, allowing bone to grow in a confluent manner, and resorbable, allowing complete replacement by new bone and restoration of near-normal host conditions. The mixing of bioactive glass particles with autologous bone is expected to provide an osteoinductive and osteoconductive material in which the osteogenic proteins derived from the host bone act with the bioactive material to promote rapid bone growth throughout the spacer used for stabilization, which must eventually depend n bony union rather than metallic fixation.

MATERIAL S Titanium surgical mesh intervertebral spacers (Depuy, Inc., Warsaw, Indiana) Particulate 45S5 Bioglassfi 100-410|im diameter. (USBiomaterials Corporation, Alachua, Florida) Autologous bone chips harvested at surgery

METHOD S Cages were placed in four adjacent vertebral (LI -L4) fenestrations created in adult sheep. Six animals were studied after three months and five after six months. The death of one animal in the second group was unrelated to the experiment. Of the four cages, one was filled with the autologous bone chips, one with Bioglass particulate and one with an approximately 50-50 mixture of the two, before insertion in the defect. The fourth cage was left unfilled as a control. The allocation of cages to sites was rotated through the animals. The cage is shown in fig. 1. In the first two animals the larger cage was used, in the remainder the cage was 10mm in diameter. Postoperativecare: Sutures were removed after 7 days. Postoperatively, the animals were monitored for signs of pain or distress until fully recovered (7-10 days). Intramuscular anti-inflammatory pain medication was administered as needed for pain. Food and water intake was monitored during the experimental period. 65

66 Bioceramics Volume 10

Fig. 1. Titanium mesh cage At autopsy the cages and surrounding tissues were removed and fixed in formalin. They were embedded in plastic and sawn sections cut transversely. The sections were stained with Paragon and Sanderson’s stains and the nature of the bone assessed. Sections with a complete and well-oriented cage cross-section were used to measure the amount of bone-fill within the cage using an Olympus Image Analysis system. The measurements were made by taking the total cross-sectional area of the cage, measured around the outer circumference and subtracting the area of the titanium mesh to give the area available for infill. The area which was not bone was then identified and subtracted from the area available and the bone infill was expressed as a percentage of the area available. In this way we hope to minimize any differences resulting from the use of 14mm cages in the first two animals and 10mm cages in the rest.

RESULT S The animals recovered well from the surgery and regained their mobility. After three months the bone infill was seen to contain particles of the Bioglassfi with normal bone (fig. 2). Space not filled by bone contained variable amounts of fibrous tissue, but no significant inflammation. After six months the appearance of the bone was similar but space not filled by bone contained little fibrous tissue and seemed to be empty.

Fig 2. Trabecular bone containing Bioglass after six months {Sanderson stain

Spinal Fusion Using TitaniumSpacers With Bioglassând AutogenousBone: J. Wilson et al. 67 Table 1 - Bone infill after 3 months Cage alone: Bioglassfi alone: Bone alone: Mixture

46.2% 56.7% 93.3% 64.2%

(20.7%-58.1%) (23.5% - 100%) (85.4% - 100%) (57.6% - 78%)

n=5 n=5 n=4 n=5

Table 2 - Bone infill after 6 months Cage alone: Bioglassfi alone: Bone alone: Mixture

47% 82.3% 98% 81.7%

(32.4% - 83%) (67.6% - 90.4%) (94.1%-100%) (69.9%-88.1%)

n=4 n=3 n=3 n=4

After three months the percentage of bone infill was as shown in Table 1. At six months however differences between test groups were less, Table 2. There was considerable variation between animals and thus a spread in the measurements but it is clear that autologous bone remains the best choice in the shorter term, although under optimal conditions complete bone fill was achieved using Bioglassfi alone in one animal. Between 3 and 6 months no change occurred in the amount of bone in the control cages and little change was possible in those filled with bone, since the fill was so high by 3 months. The infill in those filled with Bioglassfi alone and the particulate had increased to where there was no significant difference between them. The results are shown in fig. 3. Students t-test showed that there are no statistical differences between the test groups at six months.

10 0

tt

40

Cage Alone D Bioglass(R ) Alone [H Bone Alone 3 Months

(A)

Fig. 3. (A) Bone fill at 3m. (B) at 6m.

6 Months (B)

Mixture

68 Bioceramics Volume 10

CONCLUSION S We believe that we have shown in this experiment that particulate 45S5 Bioglassfi can be used to replace or dilute autologous bone used to assist in bony repair, thus reducing the immobility associated with harvesting of such bone. It appears also from these results that the contribution of autologous bone and associated osteoinductive factors may be effective even when only small amounts, such as those derived from the bleeding subchondral endplates and generated during surgery, are mixed with the Bioglassfi. The regeneration is slower than that achieved with autologous bone but the clinical advantage lies in the reduction to a minimum of the need for harvesting of bone from the patient.

POS T SCRIP T In November 1996 an 11 month-old bulldog puppy was brought to the University of Florida’s Veterinary Medical Teaching Hospital suddenly paralyzed as a result of a congenital spinal disorder, hemivertebra. She was successfully treated by stabilizing the spine with titanium mesh filled with particulate Bioglassfi and is now completely recovered and mobile. The surgeon. Dr. Roger Clemmons, explained that there is normally no treatment for this condition other than euthanasia.

ACKNOWLEDGEMENT

S

The authors thank Dr. Y. Fujishura of Tahaka University, Japan for the image analysis. The work was supported by USBiomaterials Corporation, Alachua, Florida.

REFERENCE S

1. Fraser RD, Spine,1995 20 (suppl.) pp. 167-177S. 2. Yamamuro T; Bioceramics8, 1995 Pergamon Press, Oxford England., pp.123-127. 3. Lowery GL and Harims J., Manual of InternalFixationof Spine,Raven Press, 1996 Lippencott-Raven Publishers, Philadelphia, PA, pp. 127-146. 4. Fernyhough JC, Schimandle JH, Weigel MC, Edwards CC, Levine AM, Spine,1992 17 pp. 1474-1480.

DENSE AND POROUS BIOACTIVE CERAMICS



MACROPOROU S DIPHASI C CALCIU M PHOSPHAT E CERAMICS : INFLUENC E OF MACROPOR E DL\METE R AND MACROPOROSIT Y PERCENTAG E ON BON E INGROWT H 1,2

O. Gauthier, ^J-M. Bouler, ^’Ê. Aguado, ^P. Filet and ^G. Daculsi

^ Laboratoire de Chimrgie, Ecole Nationale Veterinaire de Nantes, Route de Gachet, BP 40706, 44307 Nantes cedex 03, France, ^ Centre de Recherche sur les materiaux d’interet biologique, Faculte de Chimrgie Dentaire, 1 place Alexis Ricordeau, 44042 Nantes Cedex 01, France ABSTRAC T A total of 60 cylindrical 6 x 6 mm samples of a macroporous biphasic calcium phosphate (MBCP) ceramic was implanted into a distal femoral site in 30 rabbits. These samples represented 6 kinds of implants with 2 different macropore diameters and 3 different macroporosity percentages. Eight weeks after implantation, analysis of backscattered electron images of implant surfaces analysed by a factorial design method showed that implants with 565 îm pore size provided more abundant newly-formed bone both in peripheral and deep pores than those with 300 jam pore size. No significant differences were found between implants with 40% and 50% macroporosity, suggesting that the influence of macropore size on bone ingrowth was greater than that of macroporosity percentage. MBCP implants with 565 îm pore diameter and 40% macroporosity represented the optimal association for homogeneous and abundant bone ingrowth. KEYWORD S : bone substitute, ceramic, calcium phosphate, porosity INTRODUCTIO N Macroporosity is conducive to osteoconduction of BCP ceramics but also has many effects on their mechanical behaviour [1, 2]. Cell colonisation and bone ingrowth apparently occur when macropore size is greater than 100 |.im [3], and a reduction in macroporosity may have negative results for the biological properties of macroporous biphasic calcium phosphate (MBCP) ceramics, in that optimal macroporosity parameters have not yet been defined. The purpose of this study was to evaluate the influence of macroporosity on the osteoconduction of BCP ceramics. Bone ingrowth was quantified in several kinds of MBCP implants to determine the most desirable pore size and porosity percentage for osteoconduction. MATER[AL S AND METHOD S We used a MBCP ceramic with a 60/40 HA/pTCP weight ratio. Two main parameters were tested : macropore size (Fl) and macroporosity percentage (F2). Two different macropore diameters, 300 and 565 jam, and two different macroporosity percentages, 40% and 50%, were studied. These four values defined an experimental domain that was studied with a factorial design method (FDM), based on a first-order polynomial mathematical model, to analyse the quantitative resuhs and determine the influence of each factor on bone ingrowth [4, 5]. Ten samples each of 4 kinds of MBCP implants with different macroporosity levels (table 1) were prepared for purposes of statistical evaluation (I = 300 jam and 40% , II = 565 \xmand 40%, III = 300 |Lim and 50%, IV = 565 |im and 50%). 71

72


Table 1. Experimental matrix

1

^

1

I II III IV

Fl : macropore diameter (jLim) -1 +1 -1 +1

F2 : macroporosity percentage (%) -1 -1 +1 +1

level -1 level +1

300 –33.3 565121.7

40+1.8 50 –2.0

11-2 : interaction between Fl andF2 +1 -1 -1 +1

Ten samples of 2 other kinds of implants (V = 50% and 565 jam, VI= 30% and 300 \xm)were used on ten other rabbits to compare the experimental data for newly-formed bone with the values calculated with the FDM. Ceramic implantations (randomised distribution) were performed on 30 New Zealand white rabbits. A cylindrical defect was created at the distal end of rabbit femurs at the epiphysometaphyseal junction. A MBCP implant 6 mm long and 6 mm in diameter was positioned to fill the defect. All rabbits were killed 8 weeks after implantation. Femoral extremities were excised, fixed in glutaraldehyde solution, dehydrated in graded ethanol and embedded in glycolmethylmethacrylate. Sections of the femur from each group were analysed by undecalcified histological examination. For each sample, serial sections were cut perpendicular to the long axis of the implant. Qualitative observations were performed in light microscopy on solochrome-cyanine stained sections and with polarised light on unstained ones. The block was then sputtered with Au-Pd for scanning electron microscopy (SEM) observations. Quantitative evaluation was performed by image analysis of the SEM observations of implant surfaces using backscattered electrons (BSE). The whole surface of implants was divided in 12 contiguous fields and recorded on SEM with magnification X50. Threshold was determined by the operator on image analyser and the newlyformed bone surfaces were then automatically calculated and expressed as the percentage of the whole surface. RESULTS All implants showed extensive osteoconduction. In light microscopy, most peripheral pores were completely filled with well-mineralised lamellar bone. This new bone often showed a haversian structure. Only implants II and IV showed evidence of bone colonisation in deep pores where lamellar bone was found on the surface of almost every macropore. Measurements from SEM and image analysis observations based on graylevel distribution allowed to calculate a mean percentage of newly-formed bone for the samples of each kind of implant (table 2). According to the experimental matrix of the FDM, the influences of each factor and the interaction between Fl and F2 (I1.2) were determined for implants I, II, III and IV. The equation describing the percentage of newly-formed bone in implants can be formulated as follows : (1) newly-formed bone % = CM% + Si(fi.Fi), where CM% is the calculated mean for implants I, II, III and IV, and li(fi.Fi) is a first-order polynomial fiinction depending on significant influences and interaction between factors Fl and F2.

Macroporous Biphasic Calcium Phosphate Ceramics: O. Gauthieret al.

73

Table 2. Newly-formed bone percentages for the different kinds of MBCP implants. Implants

I

U

II I

IV

newly-formed bone % Implants newly-formed

16.7 – 3 . 9 1

20.6 – 5 . 5 0

16.8 – 3 . 0 7

22.0 – 6 . 9 3

V 22.0 – 5 . 3 0

VT 8.7 –3.28

Calculated mean (CM ) 19.0

bone % Using values from the experimental matrix, the equation became : (2) newly-formed bone % = C M % + (2.28 x F l ) + (0.39 x F2) + (0.35 x F l F2) Only the influence of macropore diameter seemed to have a significant impact on bone ingrowth. The equation can thus be simplified : (3) newly-formed bone % = C M % + 2.28 X I , XI was a coded value related to macropore diameter D by XI = (D - 432.5)/l32.5 The equation then became : (4) newly-formed bone % = 11.6 + (0.017 D) For implants V, F D M gave a percentage of newly-formed bone of 21.2% – 5.07, whereas the experimental results for 10 MBCP samples of implant V gave a percentage of newly-formed bone of 22.0% – 5 . 3 0 These two values were not significantly different. The experimental results for implants VI were not predicted by the FDM (table 2). DISCUSSIO N A FDM mathematical model was used to investigate in vivo mechanisms relative to the influence of macropore size and macroporosity percentage on bone ingrowth. Our precise experimental conditions showed that newly-formed bone in MBCP ceramics can be regarded as a simple and linear function of macropore size (11.6 + 0.017 D). FDM could not account for the behaviour of VI implants whose macroporosity parameters were chosen outside of the experimental area. This indicates that interpretation of FDM data cannot be extended beyond the experimental limits defined by implants I, II, III and IV but can predict bone ingrowth if macroporosity parameters are still chosen inside the experimental domain. Our in vivo conditions led to great variability. However, our work concerned a large number of BCP samples and provided very precise quantitative evaluation using SEM with BSE [6]. This original image analysis method based on the recording of contiguous images seems to be applicable to the study of biomaterials with good reproducibilty. Macroporosity confers osteoconductive properties on bone substitutes and it is generally admitted that 80-100 )Lim is the minimal pore size for osteoconduction. In our study, after 8 weeks of implantation, better bone ingrowth was achieved for macropores of 565 than 300 |nm. Newly-formed bone was not only significantly more abundant in BCP implants with 565 fim macropores but was also observed in both peripheral and central macropores. The BCP implants with 30% macroporosity and 300 |im pore diameter gave a very low rate of newlyformed bone. Their macropore size and macroporosity percentage were both inadequate with our

74


model. In fact, the FDM and experimental results indicated that macroporosity percentage was a less important influence than macropore size. For a similar macropore size, there was no significant difference in newly-formed bone for implants with 40% or 50% macroporosity. Our results would appear to have important mechanical implications. The mechanical properties of BCP ceramics improve in vivo due to bone ingrowth in macropores and reprecipitation of biological apatites in micropores [7, 8, 9]. In our study, a macropore size of 565 jLim provided a higher rate of newly-formed bone than one of 300 \xm.The presence of this new bone over the entire implant surface can have a favourable influence on BCP mechanical behaviour. The more macroporous implants are before implantation, the less mechanical resistance they offer [10] but it has been demonstrated that macroporosity percentage has a greater influence than macropore diameter on the compressive strength of BCP implants [5]. From this study, we could consider that implants with 565 j.im macropore size and 40% macroporosity could have a 67% higher compressive strength (24.3 MPa) compared to the same implants with 50% macroporosity (14.5 MPa). A reduction in macroporosity without major effects on bone ingrowth seems possible with 40% rather than 50% macroporosity. CONCLUSIO N This in vivo study of bone ingrowth in MBCP ceramics indicates that the influence of macropore size is greater than that of macroporosity percentage. For the same macropore size, no significant difference in newly-formed bone was noted for implants of 40% and 50% macroporosity. Osteoconduction was more efficient for MBCP implants with a 565 than a 300 jiim macropore m pore diameter and a 40%) macroporosity percentage should provide mechanical diameter. A 565 |Li improvement and preserve optimal bone ingrowth in MBCP ceramics. REFERENCE S 1. Daculsi, G. and Passuti, N., Biomaterials 1990, 11, 86-87. 2. De Groot, K., Ann.NY Acad. Sci. 1988, 253, 227-233. 3 Shimazaki, K. and Mooney, V., J. Orthop. Res. 1985, 3, 301-310. 4. Goupy, J. In: La methode des plans d’experience, Dunod, Paris, 1988. 5. Bouler, J.M., Trecant, M., Delecrin, J., Royer, J., Passuti, N. and Daculsi, G., J. Biomed. Mater. Res. 1996, 32, 603-609. 6. Skedros, J.G., Bloebaum, R.D., Bachus, K.N., Boyce, T.M. and Constantz, B., J. Biomed. Mater. Res. 1993, 27, 47-56. 7. Martin, R.B., Chapman, M.W., Holmes, R.E., Sartoris, D.J., Shors, E C , Gordon, J.E., Heitter, DO., Sharkey, N.A. and Zissimos, AG., Biomaterials 1989, 10, 481-488. 8. Daculsi, G., LeGeros, R.Z., Heughebaert, M. and Barbieux, I., Calcif Tissue Int. 1990, 46, 20-27. 9. Trecant, M., Delecrin, J., Royer, J. and Daculsi, G., Clin. Mater. 1994, 15, 233-240. 10. Le Huec, J.C, Schaeverbeke, T., Clement, D., Faber, J., Le Rebeller, A., Biomaterials 1995, 16, 113-118.


MECHANICA L FATIGU E OF HO T PRESSE D HYDROX Y APATIT E S. Raynaud, E. Champion, D. Bernache-Assollant Laboratoire de Materiaux Ceramiques et Traitements de Surface, ESA CNRS 6015, 123, Avenue Albert Thomas, 87060 Limoges, France

ABSTRAC T Polycrystalline hydroxyapatite (HAP) was densified by hot pressing. Dynamic fatigue resistance of the resulting ceramics and degradation process in aqueous solution were investigated. Inmiediate fracture strength in air decreases from 90 MPa to 40 MPa when residual porosity ratio increases from 2% to 6%. The crack propagation exponent n, characteristic of subcritical crack growth, decreases from 22.5–2 in air to 10=^4 in Ringer’s solution for materials densified at 98% of the theoretical value. A value of only n = 14–4 is obtained in air at 94% of relative density. The degradation in solution proceeds by dissolution of crystalline HAP which leads to the decohesion of grains located around residual pores at the surface of the material. KEYWORD S : Hydroxyapatite, Strength, Fatigue, Dissolution. INTRODUCTIO N Hydroxyapatite (Caio(P04)6(OH)2) is a ceramic material of interest for biological applications [1,2]. Although the mechanical properties of dense HAP, fracture strength and toughness, have been widely reported [3-7], few studies concerning mechanical fatigue are available yet [6-8]. The long time application of stresses, even at low level, can induce delayed fracture, depending on environmental conditions. For HAP, it is important to evaluate its behaviour under mechanical loads because this bioceramic is known to be chemically affected by physiological environment [9]. Fatigue phenomenon is analysed in term of subcritical crack growth with the relationship [10]: V=

da dt

= AKf A

(1)

where V is the crack velocity, a is the crack length, Ki is the stress intensity factor at the crack tip, A is a constant, and n is the propagation exponent. The value of n is characteristic of the resistance to mechanical fatigue of a material under a given environment. The fracture strength depends on the stressing rate according to the relation [11]: 75

76


Lna f

-Ln B(n + l)an-2 n+1

(2)

-Lna n+l

where a^ is the fracture strength, B is a constant, G\ is the inert fast fracture strength and d is the stressing rate. Thus, the measurement of fracture strength at different stressing rates allows the determination of the propagation exponent n from the slope of the straight line on the graph Lna^ = F(Lna). This work consisted in determining fatigue data for hot pressed HAP ceramics and evaluating the influence of the environment on materials degradation. MATERIAL S AND METHOD S A commercial stoichiometric hydroxyapatite powder was hot pressed under a constant compressive stress of 10 MPa either at 1165 C during 1 hour or at 1100 C during 30 minutes. Sintered blocks were cut into bars of 4*3*25 mm^ and each bar was polished with a 3 jim diamond paste. Quasi static, or immediate fracture strength in air was determined by three-point bending with a 16 mm span and a crosshead speed of 0.2 mm.min-l. The dissolution of HAP ceramics was investigated in Ringer’s solution at 37 C [12]. The degradation of samples was evaluated by measurements of surface roughness for immersion times ranging from one day to three weeks. Mechanical testing of HAP samples by dynamic fatigue in solution was investigated in a device which permits the control of liquid environment (constant temperature of 37 C and constant liquid flowing), the crosshead speed varied in the range 3.10-^ nmi.min-l to 2 mm.min-l. Since the determination of fracture strength measured by three-point bending can be biased by the location of the flaw which initiates crack propagation, an artificial defect was generated on the surface of samples by Vickers indentation under 4.9 N load. RESULT S AND DISCUSSIO N Immediate fracture strength in air decreases from 90 MPa to 40 MPa when the residual porosity ratio increases from 2% to 6%. Dynamic fatigue experiments allow to calculate the crack propagation exponent n (from equation 2). The results obtained from linear regressions of measured data are given in table 1. Values of n = 22.5–2 and n = 14–4 are obtained for HAP ceramics tested in ambient air and densified at 98% and 94% of the theoretical density, respectively. The propagation exponent is n = 10–2 for materials at 98% of relative density, tested in Ringer’s solution. In air, both fracture strength and resistance to subcritical crack growth decrease as the volume fraction of residual pores increases. In the same way, the liquid environment induces a drastic drop of the resistance to fatigue for HAP ceramics densified at 98%. These results are close to those found by G. De With who showed an important sensitivity of HAP Table 1. Analysis of dynamic fatigue plots. Experimental conditions HAP 98 % in air

Linear regression Lna = 4.43 + 4.23 10’^ L n a

HAP 98 % in Ringer’s solution Lna = 4.48 + 9.05 1 0 ’ L n a HAP 94 % in air HAP 94 % in Ringer’s solution

Lna = 4.04 + 6.59 lO’^Lna Not significant - too low confidence level

Meclumical Fatigue of Hot Pressed Hydroxyapatite:S. Raynaud Qi al.

77

180-

^

160140-

e

/

120-

HAP 94% HAP 98%

1 0 0^ 8 06 0-

20 30 40 50 60 70 APPLIE D STRES S (MPa)

Figure 1. Simulated lifetime versus fictive applied stress.

40^

nt-* 10 15 TIM E (days)

20

25

Figure 2. Average surface roughness versus immersion time in Ringer’s solution.

to slow crack growth in air (n = 26), mechanical degradation which is enhanced by water environment (n = 12) [7]. The results on mechanical fatigue characteristics can be clearly illustrated by the evaluation of time-to-failure under a fixed stress. Indeed, dynamic fatigue experiments allow to calculate an estimation of the lifetime of a material. The lifetime under constant applied stress is given by the following relationship : Lntf=LnBa|

n-2

- nLn Ga

(3)

where tf is the time-to-failure and a^ is the applied stress, n and B a " ^ are calculated from the linear regressions constants of dynamic fatigue data. Simulated plots of time-to-failure under constant applied stress are given in Figure 1. They show that the lifetime under mechanical loads is much shorter when the material is subjected to liquid environment. For example, in the case of HAP densified at 98 %, the lifetime under a tensile stress of 30 MPa would be of about 100 hours in solution whereas it would be of more than 100 years in air. An expected lifetime of 20 years in solution would require that stresses do not exceed 15 MPa, which means that HAP ceramics cannot be used in stressed regions of the body. This behaviour also indicates that stress enhanced chemical reaction proceeds at the crack tip, resulting in a very low resistance of HAP to subcritical crack propagation in liquid solution. The influence of liquid environment on HAP degradation can be evaluated by surface observations. Figure 2 shows the plots of average roughness (Ra) of samples surfaces versus immersion time in Ringer’s solution. The difference in initial Ra values between HAP densified at 98% and 94% is due to residual pores at the surface of materials. In any case, the average roughness increases with the immersion time in Ringer’s solution. For HAP densified at 94%, a doubling of Ra value is noticed after three weeks of immersion. A typical SEM micrograph of HAP surfaces after 3 weeks of immersion is given in figure 3a. Compared with the initial surfaces (fig. 3b), the degradation of materials surface after immersion is not uniform. Rings like grooves with dimensions close to the grains size appear in only some regions and dense regions do not seem to be degraded. It can be assessed that the degraded regions are preferentially located around residual pores presents at the surface of initial material.

78


Figure 3. SEM micrographs of HAP surfaces (material densified at 94%). (a) after 3 weeks of immersion - (b) initial. This shows that dissolution of HAP is accompanied by the decohesion of some grains. Different interpretations can be found to explain the degradation mechanism of calcium phosphates in liquid. In our case, it can be hypothesised that the degradation proceeds through HAP surface dissolution. This phenomenon would lead to a preferential decohesion of grains located around pores. CONCLUSIO N The mechanical behaviour of HAP depends strongly on the presence of residual pores. In solution, it is subjected to surface dissolution accompanied with grains decohesion around these pores. Slow crack growth is enhanced by the dissolution at the crack tip. Thus, HAP ceramics appear too brittle and sensitive to liquid environments to be used under stresses. Providing a good control of the microstructural design of HAP matrix may be obtained to prevent the detrimental effect of residual pores, composite technology seems to be a way to improve the mechanical reliability and decrease the motion of subcritical crack growth in HAP based materials and finally extend their potential applications.

REFERENCES

Oonishi, H., Biomaterials,1991, 12, 171-178. Hench, L.L., J, Amer. Ceram.Soc, 1991, 74 [7], 1487-1510. Jarcho, M., Bolen, C.H., Thomas, M.B., Bobick, J., Kay, J.F. and Doremus, R.H., J. Mater.ScLA916, 11, 2027-2035. 4. Akao, M., Aoki, H. and Kato, K., ibid.,1981, 16, 809-812. 5. Halouani, R., Bernache-Assollant, D., Champion, E. and Ababou, A., 7. Mater. Sci. Mater. Med., 1994,5,563-567. 6. Thomas, M.B., Doremus, R.H., Jarcho, M. and Salsbury, R.L., J. Mater. Sci, 1980, 15, 891-894. 7. De With, G., Van Dijk, H.J.A., Hattu., N. and Prijs, K., ibid.,1981, 16, 1592-1598. 8. Nonami, T. and Wakai, P., J. Ceram.Soc. Jpn.,1995, 103 [6], 648-652. 9. De Groot, K., In: Bioceramics,Annals New-York Acad. Sci. 1988, 227-233. 10. Evans, A.G., Int.Journ.of Fracture,1974, 10 [2], 251-259. 11. Fett, T. and Munz, D., J. Eur. Ceram.Soc, 1990, 6, 67-72. 12. Barbosa, M.A., In: Biomaterialsdegradation,edited by M. A. Barbosa, 1991, 227-252. 1. 2. 3.


MECHANICA L PROPERT Y CHANGE S IN MACROPOROU S CERAMI C AFTE R IMPLANTATIO N INT O BON E AND MUSCL E M. Trecant-Viana\ J. Delecrin\ J.M. Nguyen", J. Royer\ G. Daculsi’ ^ Centre de recherche interdisciplinaire sur les tissus calcifies et les biomateriaux, Facuhe de chirurgie dentaire, 1 place A. Ricordeau, 44042 NANTES, France ^ Unite statistique et informatique medicales, CHRU, 44042 NANTES, France Laboratoire de mecanique des structures, Ecole Centrale de Nantes, 1 rue de la Noe, 44000 NANTES, France ABSTRAC T Compressive strength and stiffness of MBCP were investigated after 1 to 18 weeks of implantation in rabbit bone and muscle. It was shown that in the two sites the mechanical properties of the implants increased with the implantation duration. Nevertheless, these changes occured to a different degree or followed a different law, suggesting site-dependant structural, physico-chemical and histological modifications. This hypothesis was confirmed by a stepwise multiple linear regression analysis relative to the two mechanical characteristics and four variables (macroporosity, bone, ceramic and microporosity). The role of newly formed bone was confirmed : it filled the macropores and confered a composite stmcture to the implant. In addition the influence of physico-chemical exchanges (dissolution/reprecipitation process) leading to a decrease of the microporosity was revealed. KEYWORD S : calcium phosphates, mechanical properties, in vivo INTRODUCTIO N Clinical applications of calcium phosphate bioceramics as bone grafts substitutes are limited by their poor mechanical properties [1-4]. Macropores are necessary to promote bone formation inside the ceramic [5], but obviously they decrease the mechanical characteristics of the implant. Yet this initial strength has been shown to change when the biomaterial was placed in contact not only with bone [6-9] but also with muscle [9]. Phenomenons taking place in implanted calcium phosphate ceramics (dissolution/reprecipitation process and bone formation) have been the subject of numerous investigations [10-19] but the mechanisms which determine the mechanical properties modifications are still unidentified. The objective of this study wâs to provide quantitative understanding of the effects of physico-chemical changes and bone formation on the compressive strength and Young modulus of implanted Macroporous Biphasic Calcium Phosphate (MBCP). MATERIAL S AND METHOD S Experiment MBCP cylinders were implanted (1, 2, 3, 6, 12, 15 and 18 weeks) into the femoral epiphysis (6x6 mm) and muscle (5x5 mm) of mature male New Zealand rabbits. Samples intended for mechanical investigation were prepared as described in a previous wôrk [8]. After compressive strength and stiffness measurements, specimens were embedded in methyl methacrylate for visual characterization of implant stmcture ; macropore, ceramic and bone 79

80


percentages were measured on microradiographs and microporosity was determined using backscattered electron (BSE) imaging [20]. Analysis The means and standard deviations for com.pressive strength and the Young modulus wêre determined and a stepwise multiple linear regression analysis wâs performed relative to the tv/o m.echanical characteristics and the four variables (macroporosity, bone, ceram.ic and microporosity) to determ,ine the best predictor of a and E of implanted MBCP. RESULT S Visual characterizatio n of implant structure Figure 1 shows the changes in bone and macroporosity after implantation into bone. It can be observed that macroporosity decreased as bone formation occured during the first 12 weeks. Then bone percentage became stable and macroporosity remained unchanged For MBCP implanted in muscle, only a slight tendancy to degradation of the ceramic was obser\êd at 18 weeks. SEM image analysis showed that microporosity decreased similarly in bone and muscle sites during the implantation period studied (Figure 2). Compressive strength and stiffness Compressive strength increased linearly in both sites during the first 15 weeks (Figure 3). Comparison of the slopes revealed that the increase in compressive strength was greater when N4BCP v/as implanted into bone (Student /-test, n=72, p’Stals precipitation [20] could have partly accounted for the increase in the mechanical strength o^ MBCP im^planted in bone.

Mechanical Property Changes in Macroporous Ceramic After Implantation:M. Trecant-Vianaet al.

81

MBC P implanted into muscle Regression analysis showed that microporosity was the best predictor of compressive strength and elastic modulus. Thus strengthening of MBCP im.planted into m.uscle appeared to be the result of apatite niicrocr>’stals precipitation in ceramic micropores. mJcroporosit y (%) 7 01

50-i

\ \ 40-]

1

H^

10

12

14

16

10

13

12

14

16

18

vwseks

Figure 1. Evolution of macroporosity ( )and bone ( ) percentages in MBCP implanted into bone.

Figure 2. Evolution of microporosity percentage of MBCP implanted into bone ( ) and muscle ( )

E (MPa)

CT (MPa)

15001 i25Cri

I

t

’ I*

^%\U. 6

8

10

12

10

14

12

14

16

18

20

w e e ks

weeks

Figure 3. Com.pressive strength of MBCP implanted into bone ( ) and muscle ( } as a fijnction of implantation duration. CT (MPa )

BoRc .site Muscle site

Figure 4. Young m.odulus of MBCP imiplanted into bone ( ) and muscle ( ) as a function of implantation duration. E (MPa) i: - 1125.3042 - 98.9118M + 34.8183C (r-0.83) f-: = 837.2275 - 60.8271K4 + 21.4962C (r=0.82)

* CT = 16.042 - 0.34534M (r=U.57)

Table 1. Results of multiple linear regression (M : Microporosity, C : Ceramic). * Simple linear regression confirming the significance of microporosity (p’, E., Lynch, K., Kerebel, B. J. Biomed.Mater. Res. 1989,23,883-894. 11. Daculsi, G., Passuti, N., Martin, S.,. Le Nihouannen, J.C, Brulliard, V., Delecrin, J., Kerebel, B. Rev. Chir. Orthop.1989, 75, 65-71. 12. Daculsi, G., Passuti, N., Martin, S., Deudon, C , LeGeros, R.Z., Raher, S. J. Z?/o/wet/. Mater Res. 1990, 24, 379-396. 13. Daculsi, G., LeGeros, R.Z., Heughebaert, M., Barbieux, 1. Ca/c// Tissue Int. 1990,46, 20-27. 14. Daculsi, G., LeGeros, R.Z., Deudon, C Scanning Microscopy. 1990, 4(2), 309-314. 15. Hardouin, P., Choppin, D., Devyver, B., Flautre, B., Blary, M.C., Guigui, P., Anselme, K. J. Mater Sci. : Mater In Med. 1992, 3, 212-218 16. Heughebaert, M., LeGeros, R.Z., Gineste, M., Guilhem, A. J. Biomed.Mater Res. 1988,22,257-268. 17 LeGeros, R.Z., Parsons, R., Daculsi, G., Driessens, F., Lee, D., Metsger, S. In: Bioceramics: Material characterizationvs. in vivo behavior, 1988, Ducheyne, P . ; Lemons, J. (eds.). New York Acad. Sci. 253, 268-271. Moore, D C , Chapman, M.W., Manske, D J. Orthop Res. 1987, 5, 356-365. 19 Passuti, N., Daculsi, G., Rogez, J.M., Martin, S., Bsamtl J.V. Clin. Orthop. andRel. Res. 1989,248, 169-176. 20 Trecant, M. Ph.D. Thesis, Nantes University, France, 1996. 21 Katz, J.L., Yoon, H.S., Lipson, S., Maharidge, R., Meunier, A., Christel, P. Calcif. TissueInt. 1984, 36, S31-S36. 22 Martin, B. Calcif. Tissue Int. 1993, 53, S34-S40. Martin, R.B., Ishida, J../ Biomech.1989, 22, 419-426


DIFFERENC E OF BONDIN G BEHAVIO R BETWEE N FOUR DIFFEREN T KIND S OF HYDROXYAPATIT E PLAT E AND BONE . S.S. Chung\ C. K. Lee^ K. S. Hong^ H. J. Yoon^ 1. Department of Orthopaedic Surgery, Samsung Medical Center. 50 ILWON-dong, Kangnam- Ku, Seoul, 135-710, Korea. 2. Department of Orthopaedic Surgery, Seoul National University Hospital. 3. School of Materials Science and Engineering, Seoul National University. ABSTRAC T The interface between four different kinds of hydroxyapatite(HAp: HA 1, HA 5, HA 6, and HA 9) and bone and the surface of the HAps were examined. The HAps were made with different starting Ca/P ratios (1.5, 1.67, and 1.83) and different maturation temperatures (30 and 90 C). Sintered HAp plates were implanted in rabbits’ tibiae, femora, and muscles of thigh. The XRD analysis, light microscopy, scanning electron microscopy, and Instron were used to examine the formation of hydroxy apatite, new bone formation, bonding behavior and tensile strength. Tensile strength was greatest between HA 9(Ca/P 1.67, 30 C) and bone, though not statistically significant. We also observed more significant new bone formation on the surface of HA 9 using light microscopy. Scanning electron microscopic examination showed partial resorption of the surface of HAp plates and mechanical as well as direct bonding between HAps and bones. [KEYWORDS : Hydroxy apatite, rabbit, tensile strength, bonding behavior] INTRODUCTIO N The autogenous bone graft has many advantages, but there are many complications and problems in harvesting autogenous bone from iliac crest [1]. The allograft as well as heterograft have many problems to be used routinely as graft material [2,3]. Hydroxyapatite has been widely studied and used in clinical field as a bone graft substitute [4,5]. There are many reports comparing biologic responses using different kinds of ceramics, such as hydroxyapatite, tricalcium phosphate, calcite, bioactive glasses, etc. [5,6]. It was our investigation that different kinds of hydroxyapatite as well as different kinds of ceramics could show different biologic responses. Four different kinds of hydroxy apatites have been made and examined for their biologic responses. MATERIAL S AND METHOD S Four kinds of hydroxyapatite powder were selected among 9 kinds of powder and compacted into plate shapes, which were then sintered at a temperature of 1300 C. The HAps were named as HA 1, HA 5, HA 6, and HA 9 and their synthetic conditions were Ca/P 1.5 (maturation temperature of 90 C), Ca/P 1.5 (30 C), Ca/P 1.83 (30 C) and Ca/P 1.67 (30 C), respectively. Eighty-four white rabbits around 3.5Kg were divided into 4 groups according to HAps used. HAp plates were inserted into proximal tibiae of all rabbits(n==21/group) through a slit of medial and lateral cortex (Fig. 1 a), into distal femora 83

84


(a) (b) Figure 1. Photos of rabbit’s tibia implanted with hydroxyapatite plate across the proximal tibial metaphysis (a) and prepared for biomechanical test (b). (n=6/group) through a slit of lateral cortex and into lateral thigh muscle (n=6/group). Seven rabbits of each group were sacrificed at 2, 4, and 8 weeks and tibiae, femora and HAps in the thigh muscles were harvested. The tibiae were wrapped with saline gauze and deeply freezed at a temperature of-70 C. One femur each sacrificed was immersed in 2.5% glutaraldehyde and the other femur of each sacrificed was preserved in formalin. HAps in the muscle was deeply freezed. After harvesting all specimens, tibiae were thawed for about 4 hours at room temperature, and then segments of proximal tibial metaphyses containing HAp plate were excised and prepared for the biomechanical test to measure the tensile strength between HAp and bone(Fig. 1 b). Traction was applied vertically to the interface between the HAp plate and the bone at a crosshead speed of 5mm/min using an Instron (Instron 8500, Instron corporation, USA). Thin segment (200 \\.mthickness) of distal femur containing HAp plate was obtained using diamond saw for the scanning electron microscopic (SEM, S2460N, Hitachi, Japan) examination to examine the interface between HAp and bone. Another thin segment of distal femur was obtained after decalcification for light microscopic observation. The surface of the HAp plates was examined with scanning electron microscopy before and after embedding of the HAp into thigh muscle. RESULT S AND DISCUSSIO N Table 1 shows average failure load between hydroxyapatite plates and rabbits’ tibiae. At two weeks after insertion of the HAp plates into rabbits’ tibiae and femora, bone and HAp did not bond together. Table 1. Failure load between hydroxyapatite plates and rabbits’ tibiae measured by Instron. HA 9 HA type HA l HA 6 HA S (1.83, 30 C) (Ca/P, temp.) (1.67, 30 C) (1.5,90 C) (1.5, 30 C) PO 2 weeks non-bonding (7) non-bonding (7) non-bonding (7) non-bonding (7) PO 4 weeks 2.54–1.48(7) 2.08 –1.04 (7) 2.25 –1.36 (7) 3.94–1.23 (7) PO 8 weeks 2.50–1.22(7) 2.12 –1.49 (7) 2.34–1.09(7) 4.01 –0.75 (7) HA : Hydroxyapatite, PO : postoperative Ca/P, temp. : Starting Ca/P ratio, maturation temperature Data : Average failure load – standard deviation (n=number of rabbits)

Bonding Behavior BetweenFour DifferentKinds of HA Plate and Bone: S. S. Chung et al.

85

(a) (b) Figure 2. Light microscopic findings 8 weeks after insertion of HA 5 (a) and HA 9 (b). Hydroxyapatite cannot be seen because of decalcification. Black dusts on (a) are remnant of hydroxy apatite. There is more prominent new bone formation on HA 9 than on HA 5. (H&E staining, x 100) At four and eight weeks, the average tensile strength was greatest between HA 9 and bone, though there was no statistically significant difference regarding the types of HAp and postoperative periods. This finding well corresponded to morphological observation using light microscopy and scanning electron microscopy. New bone formation on the HAp plates was observed on light microscopic examination and the new bone formation was more evident on HA 9 (Fig. 2 a-b). On scanning electron microscopic examination, there was very prominent new bone formation on HA 9. Irregular resorption of the surface of HAp plate and bone ingrowth into the irregularity were also more prominent on HA 9 than on other HAps (Fig. 3 a-d).

(c) (d) Figure 3. Scanning electron microscopic findings of the interface between hydroxyapatite plates and bone at 8 weeks after insertion. (x40) (a) HA 1, (b) HA 5, (c) HA 6, (d) HA 9.

Bioceramics Volume10 Di-y Pow^d-er

1300*^0, S i n t e r i n g C a / P=

1 &

.83. SO^C A

^

^ A. .

67, 30 C

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./ C a / P= 1 5, 90 C

/

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^

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(a) (b) Figure 4. XRD patterns of hydroxyapatite powder made for this study (a) reveal single phase of the powder and partial decomposition at high temperature (1300 C) to tricalcium phosphate (b). After 8 weeks in the muscle, the surface of the ceramic became granular because of partial resorption of the surface. XRD examination shows partial decomposition of HAp into tricalcium phosphate at high temperature(1300 C) and the amount of decomposition was different according to the synthetic conditions of the powder(Fig. 4 a - b). The partial resorption might reflect the partial decomposition of the HAps at high temperature, because the tricalcium phosphate was known to be subject to partial bioresorption in the biological environment [7]. This resorption resulted granular surface and this seemed to help mechanical interlocking between HAps and bone. Hydroxyapatites showed different biological responses according to the synthetic conditions. Hydroxyapatite made with Ca/P ratio 1.67 and maturation temperature 30 C showed most favorable responses in the rabbits’ tibiae and femora. Further investigation will be performed to produce porous hydroxyapatite using this biocompatible synthetic condition to fmd pore size and configuration, which can show more favorable biological responses. ACKNOWLEDGMENT The present work was supported by the grant of the ministry of Health and Welfare of Republic of Korea.

REFERENCES

1. Arrington E.D., Smith W.J., Chambers H.G., Bucknell A.L. and Davino N.A. Clin. Orthop. 1996,329,300-309. 2. Bolano L. and Kopta J.A. Orthopedics.1991, 14, 987-996. 3. Buck B.E., Malinin T.I. and Brown M.D. Clin. Orthop.1989, 240, 129-136. 4. Emery S.E., Fuller D.A., Bensusan J.S. and Stevenson S. Transactionsof the 40th annual meeting,OrthopaedicResearchSociety,1994, New Orleans, Louisiana. 156-27. 5. Jarcho M. Clin. Orthop.1981, 157, 259 - 78. 6. Neo M., Kotani S., Fujita Y., Nakamura T. and Yamamuro T. / Biomed.Mater.Res. 1992, 26, 255-267. 7. Renooij W., Hoogendoom A., Visser W.J., Lentferink R.H.F., Schmitz M.G.J., van leperen H., Oldenburg S.J., Janssen W.M., Akkermans L.M.A. and Wittebol P. Clin. Orthop.1985,197,272-285.


TREATMEN T OF OSTEOMYELITI GLAS S CERAMI C BLOC K

S BY ANTIBIOTIC-SOAKE

D POROU S A-W

K. Kawanabe, Y. Okada, H. lida, and T. Nakamura Department of Orthopedic Surgeiy, Faculty of Medicine, Kyoto University, 54 Shogoin-Kawaharacho, Sakyo-ku, Kyoto 606, Japan. ABSTRAC T A new dnig delivery system was developedforosteomyelitis using apatite-wollastonite containing glass ceramic (A-W GC) that had been soaked with antibiotics under high vacuum. An 8-mm^ porous A-W GC block (porosity; 70% and 20-30%) and hydroxyapatite (HA) block (porosity; 35-48%) were placed in a bone cement mixer, and mixed with an antibiotic solution. The slow release activity of two antibiotics, isepamicin sulfite and cefinetazole sodium,fromthe porous blocks was tested. An evaluation was made of the slow-release capabilities of the isepamicin sulfitefromthe porous A-W GC block (porosity; 70%) which was maintained at more than 0.5 îg/ml after 28 days. However, that from the porous HA block was less than 0.5 ng /ml after 14 days. In a clinical study, two patients with osteomyelitis, including one with infected hip arthroplasty and osteomyelitis of the tibia, were treated and thefecihad completely healed by the end of the follow-up period. INTRODUCTIO N Chronic osteomyelitis is difficult to treat due to the characteristics ofbone, and the object of treatment is to maintain the bactericidal concentration of antibiotic at the infection focus long enough for the healing process to begin. Various antibiotic carrier systems have been developed, including one in which bone cement is mixed with antibiotic-impregnated polymethylmethacrylate (PMMA) beads [1.2] . However, the problem with the use of PMMA beads inserted locally is that subsequent surgery is required for replacement with an autograft. Recently, dmg delivery systems (DDSs) using resorbable materials, collagen [3] , fibrinogen [4] and polylactic acid [5] have been developed. Although it is not necessary to remove them, they cannot be used tofillthe infection site with new bone without bone grafting. We have developed a new DDS using antibiotic-soaked porous A-W GC block, which was demonstrated previously to forma chemical bond with living bone and to have a mechanical strength nearly equal to that of cancellous bone [6] . MATERIAL S AND METHOD S In vitrostudy Two types of porous apatite-wollastonite containing glass ceramic (A-W GC: Nippon Electric Glass Co., Ltd, Otsu, Jân) werefibricatedin 8-mm^ blocks of porosity is 70% (A-W GC 70) and 2030% (A-W GC 20-30), with pore sizes of 200 ^m and 10-50 ^m, respectively. A porous hydroxyâtite (HA) block of the same size, porosity 3548%, pore size 50-300 [im (Bioceram: Sumitomo Pharmaceutical Co.. Ltd, Tokyo, Jân) was used as a control (Fig.l). Two kinds of antibiotic, isepamicin sulfate (ISP: C22H43N5O12 x H2SO4, (X ^2), MW: 569.61) andcefmetazole 87


îia.-a

liiiiiili^f c

1 ^îs^^W-^m^’W^MmMmB^, (a) (b) Figure 1. SEM appearance of (a) surface of A-W GC ( porosity 70%) and (b) HA (porosity 35-48%) (CMZ: Ci5Hi6N705S3Na, MW: 493.51) were used. An experimental study of DDS with antibiotic-soaked porous blocks was carried out follows. The three kinds of porous ceramic block were placed in a bone cement mixer (Mixevac 11 High Vacuum System, Stiyker, MI, USA) and mixed with solutions of the two antibiotics, ISP and CMZ (100 mg/ml), and vacuumed at about 500 mmHg for 10 min to allow the antibiotics to soak into the pores. The antibiotic absorption rates of A-W GC 70, A-W GC 20-30 and HA determined by this method were 76.1%, 21.8% and 25.3%/ 8 mm^ volume, respectively. To estimate the concentration of slow-release antibiotic, the blocks were stored in PBS (3 ml) at 37 C. andthePBS was replaced every two days. Preserved PBS containing the released antibiotic was stored at -20t;. An in vitro elution study was then performed using antibiotic assay by high-perfonnance liquid chiomatogi^hy (HPLC), Clinical case Two patients were treated using this method. A 35.year-old man was operated on for osteomyelitis of the right proximal tibia Abscessformationwas observêd in the sameregionnine years after the primary operation, and Salmonella was cultured from this specimen. Afier undergoing curettage, appropriate A-W GC blocks were placed in a cement vacuum mixer and soaked with the antibiotics CMZ and ISP. The A-W GC blocks were trimmed and inserted into the osteomyelitis focus. The other patient was a 55-year-old hemophilia man, who was sufieringfix)minfeaed arthroplasty. He underwent revision surgery using antibiotic-soaked A-W GC blocks. The follow-up terms were lyr 6mo and lyr, respectively. RESULT S An evaluation was made of the slow-release capabilities of ISP, and the level was maintained at at least 0.5 ng/ml after 28 days in both A-W GC 70 and 20-30 blocks. However, the level from the porous HA block was less than 0.5 îg/ml after 28 days. In the case of CMZ, the three kinds of porous block showed a level of less thanO.5 [ig/ml even after 14 days (Fig. 2). The mean release ratio (antibiotic released in PBS / antibiotic soaked in block) of A-W GC 70, A-W GC 20-30 and HA were 91%, 100% and 100% for ISP, and 48.7%, 37.9% and 42.3% for CMZ, respectively.

Treatmentof Osteomyelitisby Antibiotic Soaked Porous A-W Glass Ceramic: K. Kawanabe et al. CMZ(ug/mi)

ISP(ug/ml) 1 0 0 0 0 0 (] D A-W GC (70% )

D A-W GC (70% ) @ A-W GC (20-30% )

0 A-W GC (20-30% )

HA

2 d 4 d 6 c l 8 d

14 d

18 d

HA

0,1

L

2d 4 d 6 d 8d

14 d

(a) (b) Figure 2. A gr^h showing therateofreleaseof ISP (a) and CMZ (b)fromA-W GC 70, A-W GC 20-30 and HA. hi the dinical cases, both the feci had healed at the end of the follow-up period without complications. The border between the A-W GC blocks and bone became unclear lyr 6mo afer surgery in the case of osteomyelitis in the right proximal tibia (Fig. 3).

(a) (b) Figure 3. A 3 5-year-old-man, osteomyehtis of the right proximal tibia was recurred after nine years afier primaiy operation, (a) Radiogr^hs made one week after second operation. After curettage the infection focus, porous A-W GC blocks were soaked with CMZ and ISP, and implanted, (b) Radiograph made lyr 6mo after second operation. The border between the porous A-W GC and bone became unclear compared with one week.

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DISCUSSIO N Antibiotic-loaded acrylic bone cement beads has been studied in detail and used clinically. However, it must be removed by a further operation, and its long-term implantation is difficult. Some authors have reported that PMM A bone cement has a disadvantage of thermal damage to the antibiotic [1.2] Recently, biodegradable materials have been developed as DDSs for antibiotics. However, these materials can not fill the dead space and may be afocusforrecurrent infection over a long period. Bioactive ceramic is an ideal DDS fiom this view point. Although HA was used as a DDS for antibiotics in a several reports [7.8] , A-W GC has been demonstrated to have higher mechanical strength and bioactivity than HA, and A-W GC 70 absorbed andreleasedmore antibiotic than the porous HA block during a one month period in this study. The high-vacuum system using a cement mixer was efective for soaking the antibiotics into the ceramic pores REFERENCE S 1. Whaling, H., Dingelden, E., Bergmann, R., Reuss, K. The release of gentamicine fix)m polymethylmethacrylate beads. J. Bone Joint Surg., 1975. 60-B. 270-275. 2. Baker, AS., Greenham, L.W., Release ofgentamicinfix)macrylic bone cement. J. Bone Joint Surg., 1988. 70-A.1551-1557. 3. Ascherl, R., Stemberger, A., Lechner, F. Behandelung der chronischen osteomyelitis mit einem koUagen-antibiotika-verbund-vorlaufrge mitteilung. Umfall Chirurg., 1986. 12. 125-127. 4. Zilch, H., Lambiris, E. The sustained release ofcefotaxin fiom afibrin-cefotaxincompound in treatment of osteitis. Arch Orthop Trauma Surg. 1986. 106. 36-41. 5. Wei, G., Kotoura, Y., Oka, M., Yamamuro, T., Wada, R., Hyon, S.H., Ikada, Y., A bioabsorbable deliveriy systemforantibiotic treatment of osteomyelitis. J Bone Joint Surg., 1991. 73-B. 246-252. 6. Nakamura, T., Yamamuro, T., Higashi, S., Kokubo, T., Itoo, S. A new glass-ceramicforbone replacement: Evaluation of its bonding to bone tissue. J Biomed Mater Res., 1985. 19. 71-84. 7. Shinto, Y., Uchida, A.,Korkusuz, F., Araki, N., Ono,K. Calcium hydroxyapatite ceramicused as a delivery system for antibiotics. J Bone Joint Surg., 1992. 74-B. 600-604. 8. Itokazu, M., Matsunaga, T., Kumazawa, S., Oka, M. Treatment of osteomyelitis by antibiotic impregnated porous hydroxyapatite block. Clin. Mater., 1994. 17. 173-179.


CALCIU M HYDROXYAPATIT E CERAMI C IMPLANT S IMPREGNATE D WIT H ANTIBIOTI C FOR THE TREATMEN T OF CHRONI C OSTEOMYELITI S YasuoYamashita,Toru Yamakawa ,Kou Kato Yoshitaka Shinto*, Nobuhito Araki*, Atsumasa Uchida Department of Orthopaedic Surgery, Mie University Faculty of Medicine,Ed)basi 2-174, Tsu-shi, Mie-ken 514, Japan *Department of Orthopaedic Surgery, Osaka University Medical School, Yamada-oka, suita-shi 2-2, Osaka-Fu, 565, Japan. ABSTRAC T Twenty patients with chronic osteomyelitis were treated by implanting calcium hydroxyapatite ceramic with antibiotic into a cavity produced after through surgical excision of necrotic tissue. Within 3 months all of the infected sites had healed. During the period of follow-up ranging from 3 to 75 months we have never experienced a recurrence of infection. There were 3 of those patients had infected prostheses and were successfully revised One patients underwent one stage revision surgery, and another two patients underwent two stage operation. Not only was infection controlled, but there was incorporation of the ceramic material into host bone as judged by radiography. We recommend the use of porous pieces of calcium hydroxyapatite impregnated with antibiotic as a new drug delivery system for the treatment of chronic osteomyelitis. KE Y W O R D S Antibiotics, Hydoroxyapatite, Drug delivery system. Osteomyelitis INTRODUCTIO N Chronic osteomyelitis is known to have difficult surgical problem, particularly in the developing world, despite advances in surgery and more than fifty years experience with antibiotic therapy. Two principles of treatment are paramount: necrotic tissue which has a blood supply unnable to promote normal healing process must be removed, and appropriate antibiotic drugs must be administered [1]. Porous calcium hydroxyapatite (CHA) which is similar to bone mineral composition has excellent biocompatibility, can resist mechanical forces, and is effective in filling cavities and defects in bone [2]. We have already reported porous CHA is very effective as a slow release system for antibiotics in an animal model. [3,4] We have now used porous CHA impregnated with antibiotic clinically and report our experiences, and we believe that this new system is simple, can be performed safely in some few stage, and offers satisfactory results. MATERIAL S AN D METHOD S We treated 11 men and 9 women with chronic osteomyelitis using the principles of surgical debridement, local implantation of CHA impregnated with antibiotic, and systemic 91

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antibiotic administration. The mean age of the patients at the time of treatment was 39.1 years (range: 14 to 77 ). The duration of disease was judged to be from 12 to 132 months (mean: 33 months). At the time of initial presentation all patients had clinical andôr radiological evidence of chronic osteomyelitis. The chronic infection occurred after acute hematogenous osteomyelitis in 12 patients, following open fracture in 3, after closed fracture in 1, and after joint replacement in 3. Nine patients had one or more draining sinuses. Each site of infection was initially aspirated in order to detect a causative organism. The pathogens cultivated were Staphylococcus aureus in 8 cases, Staphylococcus epidermidis in 3, Pseudomonas aeruginosa in two, and Streptococcus pyogenes. Streptococcus pneumoniae, and Klebsiella pneumoniae in each one. The choice of antibiotic for impregnation into the CHA ceramic material was determined by the sensitivity of the cultured organism to drugs. In the 4 patients in whom no organisms were grown broad spectrum antibiotics were selected. We assessed healing by the clinical picture, laboratory findings, and radiological evidence of incorporation of the CHA implant and remodeling of surrounding bone. The duration of follow-up was from 3 to 75 months (average: 47.9 months). Preparation of CH A ceramic impregnated with antibiotic. CHA ceramic blocks were sintered at 1200 C for two hours and had a porosity of 30% to 40% with diameter of the micropores between 40 and 150 micrometers. There was an interconnecting pore structure open to the external surface of the blocks. Operative procedures. The bone cortex was fenestrated to a size permitting removal of all necrotic bone, sequestra, and pathological granulation tissue. During the necrectomy, the chosen antibiotic was packed into a central cylindrical cavity in each porous block and the cavity then sealed with a CHA plug. (Figure 1) The volume of antibiotics packed into the cavity depended on the size of the cavity within the different blocks. The usual range of antibiotic dose in each ceramic block was between 100 and 400 mg. The antibiotics were used, either alone, or in combination. The excavated defect in the bone was then packed with the CHA ceramic pieces which had been each impregnated with the chosen antibiotic. Various sizes and number of ceramic block were used so that the excavated bone defect could be completely filled.

Figure 1 Illustration of CHA impregnated with antibiotics.

Figure 2 Case 1 Chronic osteomyelitis of the proximal region of the tibia.

Calcium Hydroxyapatite Ceramic Implants Impregnated With Antibiotic: Y. Yamashitaet al.

93

CASE R E P O R T S Case 1 An 18-year old man had complained of dull pain in the proximal leg and around the knee. The radiographs showed a sclerotic thickening of the cortex in the proximal tibia. The lesion was opened, anddebrided, and then packed with number of CHA ceramic blocks which had been impregnated with fosfomycin sodium anddbekacin sulfate, ( as no causative organism was cultured ).Four months after the operation, the lesion had been completely healed (Figure 2). Case 2 A 40-year old man with osteosarcoma of the distal femur had infection after wide resection with reconstruction by a tumor knee prothesis. The prosthesis was removed and pathological granulation tissues were debrided and the antibiotic-impregnated CHA blocks were placed in dead space. After 3 months the revision surgery with another tumor knee prosthesis was performed The patients had no recurrence of infection and maintain excellent function after 3 years (Figure 3). Case 3 A 67-year old woman had infection of a knee prosthesis inserted for the treatment of osteoarthritis. She was treated firstly with antibiotic-impregnated Polymethylmetacrylate (PMMA) beads. She had persistent pain and swelling of the knee. A biopsy indicated that Staphylococcus epidermidis was the pathogenic organism. The prosthesis was removed and all necrotic tissue were carefully debrided The antibiotic-impregnated CHA ceramic blocks and another prosthesis was inserted in one stage. Twelve weeks after the revision surgery there had been no recurrence of infection (Figure 4).

Figure 3 Case 2 Chronic osteomyelitis of the mega prosthesis for osteosarcoma in the treatment of distal femur.

Figure 4 Case3 Inplantation of CFIA drug delivery system for the infectedTKA.

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RESULTS At the latest follow-up examination for each of the 20 patients all foci were completely healed Fifteen had pain relief and improvement of laboratory abnormalities within 4 weeks after surgery. In the rest there was resolution of infection within 12 weeks. Recurrence of infection has never occurred after this treatment for follow-up period Radiolucent zones around the ceramic implant gradually disappeared over six months, and in some cases homogenous intramedullary radiodensity surrounded the CHA blocks. There was no radiological evidence of degradation of the ceramic, but in those patients in whom we performed ’second look’ surgery after healing there was some histological evidence of ceramic degradation. DISCUSSIO N It is essential to maintain a high concentration of an appropriate drug at the affected sites for a sufficiently long time, in order to obtain complete eradication of infection of bone and soft tissues. Because of the altered structure of the tissues surrounding an infected site the diffusion of antibiotic drugs into the central part of the infection may require high serum concentration of the drugs. This may cause side effects such as myelosuppression, renal failure, and hepatitis. It is possible to increase the local concentration of antibiotics by impregnating them into carrier vehicles which are implanted into the infected site. PMMA used generally as a bone cement has been the most widely evaluated[5,6,7]. The disadvantages include reduced biocompatibility with bone, short duration of antibiotic release, very low release rate, thermal damage to the antibiotic, and the requirement to remove the PMMA at the end of therapy. Nevertheless this method has been widely used for the surgical treatment of chronic osteomyelitis. Drug delivery system with porous hydroxyapatite ceramic may be effective to apply an appropriate drug for various di sease such as pyogenic osteomyelis, tuberculous osteomyelitis. We are of the opinion that antibiotic-impregnated CHA ceramic is superior to acrylic bone cement systems. Many antibiotic can be placed in a CHA as there is no thermal damage to the drug. All of the impregnated antibiotic can be released over a long period and none is trapped in the ceramic. Biomechanical properties of CHA is similar to those of bone, and the composite of ceramic with newly-formed ingrowth of bone into the pore is almost same as the original bone. As a consequence the antibiotic-CHA ceramic composites both control the infection, restore mechanical strength, encourage osteoconduction into their pores, and avoid the need for further surgery. From these findings, we believe that this new system is simple, can be performed safely in one stage, and offers satisfactory results.

REFERENCES

1. Norden,C.W., Gillespie,W. J. and Nade,S. Infectionin Bones andJoints Boston,Blackwell, 1994, 3-418. 2. Uchida,A., Araki,N., Shinto,Y., Yoshikavva,H., Ono,K. and Kurisaki,E., J .BoneJoint Surg [Br] 1990, 72-B, 298-302. 3. Shinto,Y., Uchida,A., KorukusuzJF., Araki,N. andOno,K., yôn^/om/^wrg 1992, 74-B, 600-604. 4. KorukusuzJF., Uchida,A., Inoue,K., Shinto,Y., Araki,N. andOno,K. J Bone Joint Surg [Br] 1993, 75-B,111-114. 5. Buchholz,H.W., Elson,R.A., andHeinert,K. Clin Orthop1984, 190,96-108. 6. Bayston,R., and Milner,R.D. J Bone Joint Surg [Br] 1982; 64-B, 460-464. 7. Boda,R. Arc/i OrthopTraumaSurg 1982, 101:39-45.

BONE CELLS ONTO BIOACTIVE CERAMICS



MEASUREMEN T OF INTAC T OSTEOCALCI N CONTENT S IN THE COMPOSIT E OF POROU S HYDROXUAPATIT E CERAMI C AND ALLOGENEI C MARRO W CELL S M. Akahane, H. Ohgushi, T. Yoshikawa, S. Tamai, Y. Dohi,* K. Hosoda** and T. Ohta** Department of Orthopedics, and Public health*, Nara Medical University, Kashihara city, Nara 634, Japan ; Teijin Institute for Bio-Medical Reserch**, Teijin Ltd., Hino city, Tokyo 191, Japan

ABSTRAC T Osteocalcin is synthesized particularly by osteoblast as an extracellular matrix protein. We measured intact osteocalcin contents in allogeneic rat marrow cells/hydroxyapatite (HA) composites implanted at rat subcutaneous sites. At 4 weeks after implantation, bone formation was not detected and only a trace of the osteocalcin was detected in the composite. However, under the immunosupression with FK506, bone formation together with abundant osteocalcin was detected in the composite, and the osteocalcin content was comparable to that of isogenic marrow/HA composites. These results indicate that under the immunosupression, allogeneic bone marrow cells can differentiate into active osteoblasts of which activity is comparable to that of isogenic marrow cells. KEYWORD S Osteocalcin, Hydroxy apatite.

Allogeneic bone marrow

INTRODUCTIO N We have reported that subcutaneous implantation of HA ceramics combined with marrow cells show new bone formation [1]. Osteocalcin (bone Gla protein) is a major noncoUageneous protein in bone matrix and exclusively synthesized by osteoblast. Biochemical analysis of the marrow/HA composites showed that the osteocalcin begun to appear at about 3 weeks when the obvious bone formation initiated, then the osteocalcin contents and bone area increased as time passed [2]. Therefore, the osteocalcin is a useftil biochemical parameter to identify bone tissue. For measuring the osteocalcin, the harvested composites were immediately crushed, homogenized and then measured by using radioimmunoassay (RIA). For some cases, the ceramics were frozen until the assay of osteocalcin [2,3]. Recently, we reported that not only isogenic cells but allogeneic [4,5] cells show bone formation under immunosuppression with FK506. We also established the method of measuring intact rat osteocalcin using anti N- and anti C-terminal rat osteocalcin antibodies raised against a Nterminal 20 residues peptide and a C-temiinal 10 residues peptide of rat osteocalcin [6]. In 97

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this experiment, we measured intact osteocalcin in the cryopreserved composite which was combined with either allogeneic or isogenic marrow cells, and focused on the effect of freezing and immunologic barrier. MATERIAL S AND METHOD S Marrow cellpreparationand Implantationof ceramic Male 6-week-old ACI and 7-week-old Fischer 344 rats were used for donor. Syngeneic 7week-old Fischer rats were used for recipient. The magnitude of the immunological mismatch between ACI and Fischer rat is major. Marrow cell suspensions (5 x 10 ^ nucleated cells/ml) from ACI and Fischer rats were prepared as reported previously [4]. Disk shaped (5mm diameter, 2mm thickness) HA ceramics (Interpore 200, Interpore International, Irvine, Cahfomia) were immersed in each cell suspension from ACI or Fischer rats, then implanted on the back of a recipient Fischer rat. AdministrationofFK506 and Harvest of ceramic FK506 (Fujisawa Pharmaceutical Co., Ltd., Osaka, Japan) was suspended in saline and administered intramuscularly to the recipient rats (Img/Kg/day). As a control, saline was administered. All rats received FK506 or saline every day for 2 weeks and every second days for additional 2 weeks. The ceramics were harvested at 4 weeks after the implantation and stored at -80 C for about 1 month.

Peroxidase | anti N-20 Ab ^^/^N^

N terminal

I

20 peptide s

10 peptide s

1

3

C terminal

^N"" ^

anti C-10 Ab

Figure 1. Schema of the intact osteocalcin measurement by sandwich immunoassay.

IntactOsteocalcinContentsin theCompositeof HA and AllogeneicMarrow Cells: M. Akahaneet al. 99 Measurement of intact osteocalcin

The frozen ceramics were crushed, homogenized in 0.2 % Nonidet P40 containing 1 mM MgCl2 and centrifuged. Osteocalcin was extracted from the sediment by shaking in 2 ml of 20 % fonnic acid for 2 weeks at 4 C. An aliquot (500 \i\) of the fonnic acid extract was then applied to a colunm of Sephadex G-25 and eluted with 10 % fonnic acid. Protein fractions were collected, lyophilized and used for measurement of intact osteocalcin. The principle of the measurement is based on the sandwich iminunoassay which recognizes both N and C tenninal peptides of osteocalcin molecule (Fig. 1). A peroxidase conjugated with a rabbit F (ab’)2 fragment of the anti-N-tenninal 20 residues peptide antiserum (anti-N-20) was prepared [7], Polystyrene balls were dipped in anti-C-10 IgG in phosphate buffered saline (PBS) and incubated at 4 C overnight. Immobilization was tenninated by rinsing the balls with PBS, followed by coating with 1 % bovine serum albumin-PBS at 4 C for 2 days. Standard solutions of purified rat osteocalcin were prepared at concentrations of 0-5 ng/ml. Two hundreds fil of standard solution and 200 \A of peroxidase-labeled anti-N-20 IgG solution with the anti-C-10 IgG-fixed balls were placed in glass tubes. After incubation for 1.5 h at 37 C, each ball was washed three times with saline, then 0.4 ml of tetramethylbenzidine and 0.017 % hydrogen peroxide were added to the tube. The mixture was incubated at 37 C for 30 minutes and the enzyme substrate reaction was tenninated by adding 1 ml of IN H2SO4. The enzyme reaction product was measured by the absorbance at 450 nm. RESULT S AND DISCUSSIO N In this experiment, the harvested ceramics (marrow/HA composites) were immediately immersed into liquid nitrogen and stored at -80 C for about 1 month. Then the ceramics were crushed and maintained at 4 C to extract osteocalcin for about 2 weeks in 20% formic acid. As shown in Table 1 (without FK506), mean intact osteocalcin content in the frozen and stored isogenic marrow/HA composite was 0.68 jug/implant. The content was comparable to 0.59(ag/implant in non-frozen composite which was immediately crushed at the time of harvesting and followed by osteocalcin extraction in 20% formic acid. The osteocalcin

Table 1. Bone fonnation and osteocalcin contents (|ig/implant) in marrow/HA composite, (data are mean – SEM).

Allografts^ Isografts ^ Isografts ^ 1) 2)

With FK506 Osteocalcin Bone fonnation contents + 1.017 –0.224 + 0.854 –0.179 "

" - -

-

-

_

_

Without FK506 Bone fonnation Osteocalcin contents 0.036 –0.004 + 0.682 –0.210 + 0.588 –0.165

The data show the intact osteocalcin contents measured by sandwich immunoassay as described in Materials and Methods. The osteocalcin contents in the isografts (isogenic marrow/HA composites) were measured by conventional RIA as described in ref [3]. The composites were crushed immediately after harvesting and followed by osteocalcin extraction in formic acid.

100


contents of the non-frozen composites were detennined by conventional radio-immunoassay (RIA). The data indicate that there was few degradation of osteocalcin molecule during the steps of the measurement and therefore the molecule is quite stable under the low temperature of -80 C and 4 C in the presence of forniic acid. Furthennore, the data indicate the reliability of our previous reports of osteocalcin measurement (RIA) in marrow/HA composite. The bone formation occurred in allogeneic m a r r o w / H A c o m p o s i t e under immunosuppression with FK506. However, it was not observed in allogeneic composite without FK506 and only a trace of the osteocalcin was detected in the composite (Table 1). In this experiment, we measured intact osteocalcin molecule, because it is known that degradation of bone tissue accompanies the degradation of osteocalcin molecule. Therefore, measuring the intact form is crucial in identifying nonnal bone tissue. As shown in Table 1, the amount of osteocalcin in the allogeneic marrow/HA composite with FK 506 was comparable to that of the isografts (with and without FK506). The data of this experiment i n d i c a t e that the bone formed in allogeneic m a r r o w / H A c o m p o s i t e s u n d e r immunosuppression with FK506 did not show rapid degradation which might be initiated by the immunological reaction. Therefore, the surface of HA can support natural process of osteoblastic differentiation of allogenic marrow cells under immunosuppression with FK506.

REFERENCES 1. 2. 3. 4. 5. 6. 7.

Ohgushi, H., Goldberg, V. M. and Caplan, A. I J.Ortop.Res.,7:568-578,1989. Yoshikawa, T., Ohgushi, H., Okumura, M., Tamai, S., Dohi, Y. and Moriyama, T. Calcif Tissue Int., 50:184-188,1992. Inoue, K., Ohgushi, H., Yoshikawa, T., Tamai,.S., Dohi, Y., Hosoda, K. and Ohta, T. Bioceramics Volume. 8:99-102,1995. Sempuku, T., Ohgushi, H., Okumura, M. and Tamai, S. J.Orthop.Res.,14:907-913,1996. Sempuku, T., Ohgushi, H., Okumura, M. and Tamai, S. Bioceramics Volume. 8:397401,1995. Ohta, T., Azuma, Y., Kiyoki, M., Eguchi, H., Hosoda, K., Tsukamoto, Y. and Nakamura, T. Calcif Tissue Int., 59:283-290,1996. Fujiwara, K. Yasuno, M, and Kitagawa, T. Cancer Res., 41:4121-4126,1981.

Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proeeedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd

SI-CA- P XEROGEL S AND BON E MORPHOGENETI C PROTEI N AC T SYNERGISTICALL Y ON RA T STROMA L MARRO W CEL L DIFFERENTIATIO N IN VITRO E. M. Santos^, P. Ducheyne^’^, S. Radin^. B. Shenker^, I. and Shapiro^ Departments of ^Bioengineering, ^Pathology, ^Biochemistry and Ôrthq)aedic Surgery, University of Pennsylvania, Kiiladelphia, PA 19104.

ABSTRAC T The effect of a novel bioactive xerogel glass carrier with and without bone morphogenetic protein (BMP) on the osteogenic activity of rat stromal marrow cells was studied in vitro. Cell differentiation was more pronounced on xerogel glass without BMP than that of cells grown on plastic with BMP. Stromal cell differentiation, as measured by alkaline phosphatase activity and osteocalcin synthesis was most increased when the BMP was incorporated or adsorbed onto the xerogel glass. The data suggest that the xerogel glass concentrates osteoinductive proteins at its surface and potentiates their function. KEYWORDS : bioactivity, growth factor, cell culture, cell differentiation INTRODUCTIO N Fracture non-unions and large bone defects represent major clinical pDblems in the practice of reconstructive orthopaedic surgery. ^ Since current treatments for these conditions, such as autogenous bone grafting, have limitations inherent in their use, new approaches for bone tissue repair are valuable.^’*’ One novel approach is the use of osteoinductive bone growth factors, such as bone mOTphogenetic proteins (BMP).^ Bioactive glass has been shown in numerous studies to bond to bone in ydvo. In our group we have shown that porous bioactive glass can serve as an effective template for the growth of bone like tissue in vitro.^These studies also revealed the importance of pre-treating the glass surface. The treatment led to the formation of a calcium phosphate surface layer with proteins adsorbed and incorporated into it. With this treatment neonatal rat calvaria osteoblasts expressed the markers of the osteoblast phenotype extensively within 4-7 days of culture. In contrast, without the treatment, the osteoblast phenotype was not yet expressed within the same culture duration. Using sol gel synthesis exclusively at room temperature, a glass has been made that releases functional bone growth factors in a sustained manner over a period of several weeks. ^ In this paper we document the effect of this material with and without BMP-2 on the proliferation, differentiation and function of rat stromal marrow cells. 101

102


MATERIA L AND METHOD S Synthesis: Xerogel discs with a composition of 70% Si02 - 25% CaO - 5% P2O5 (S70) were synthesized using a room temperature sol-gel procedure. Tetramethylorthosilane, calcium metbov^’^thoxide and triethylphosphate were mixed under an argon atmosphere. After casting, soluuu.. ryith or without BMP in 0.1 N acetic acid mixed with a-MEM containing 15% fetal bovine serum was added to sol samples. Solutions with BMP (recombinant human BMP-2, Genetics Institute, Cambridge MA) contained 25 ^lg of it. Gels were aged and dried to 50% of their original weight. The resulting discs (10 mm in diameter and 4 mm in height) were sterilized by exposure to ultraviolet light. Material pre-treatment to form a calcium phosphate surface laver was conducted before cell culturing by immersion in sterile Dulbecco’s phosphate-buffered saline (GibcoBRL/Life Technologies, Grand Island, NY) for 3 hours. Treamient parameters were selected such that the treatment would not cause a significant loss of incorpwated protein from the sol-gel prior to the cell culture experiments. After the treatment, Fourier transform infrared spectroscopy (FTIR) (Nicolet 5DXC) was performed to establish that P-0 bend peaks were present, thereby revealing the formation of calcium phosphate layer. As a second pre-treatment step, for which we developed the rationale in our lab before"*, S70 discs were immersed in 3 ml of tissue culture medium (TCM) containing a-MEM + 15% fetal bovine serum for 1 hour prior to cell culture. Rat Stromal mam?W ggUs were harvested from 4-5 week female Wistar rats using the methods described by Maniatopoulos et al.^ Isolated cells were plated on tissue culture plates in medium containing a-MEM + 15% FBS with 50 U/ml penicillin, 50 ^g/ml streptomycin, and 10" 8 M dexamethasone in a 37 ""C, 5 % CO2 - 95% air incubator. Non-adherent cells were removed by washing after 24 hours. Thereafter, the medium was supplemented with 50 M^g/ml of ascorbate and exchanged every two days. Once the cells were confluent (after 1 week of primary culture) the adherent cells were detached using 0.25% trypsin in Ca- and Mg- free Hank’s Balanced Salt Solution and resuspended in culture medium. 1 x 10^ stromal cells in a 100 |xl TCM solution were seeded on the surface of S70 discs or tissue culture plates (35 mm in diameter) and allowed to attach for one hour. Medium was then added to the culture dish and incubated for either 6 or 10 days. The medium was exchanged every other day. Control groups included cells cultured in tissue culture dishes without BMP (C) otwith 10 ng BMP added to the initial medium and with every medium exchange (C-BT). Experimental groups included sol-gel discs without BMP (SG), or with BMP added as follows: 25 ^g of BMP incorporated into the S70 discs (SG-BI) , 100 ng of BMP added to TCM during the second pre-treatment step (SG-BP), 10 ng of BMP added to the initial medium as well as with every medium exchange (SG-BT) . Aliquots of the medium were collected before every exchange and before cell harvesting. Cell lysate was obtained by aspirating the TCM from the plates, washing the plates with PBS, and then extracting the sample with 1 ml of 3% Triton X-100 in PBS. Cell extracts were analvzed for total protein content, total DNA content, alkaline phosphatase activity (AP) and osteocalcin production using techniques described elsewhere.^ AP activity and osteocalcin synthesis results were normalized to cell number (DNA content) and surface area available for cell growth. Collagen typing was performed by SDS-polyacrylamide gel electrophoresis (SDS-PAGE) run at 100 mV.

Si-Ca-P Xerogels and BMP on Rat Stromal Marrow Cell DifferentiationIn Vitro: E.M. Santos et al.

103

RESULT S The average DNA content of groups and alkaline phosphatase (AP) activity, normalized by DN A content and surface area available for cell culture, are displayed in Figures 1 a, b.

mc 0 0 ^

m 3 days

10 days

1

SG

SG-BI SG-BP SG-BT C-BT

10 days

Figure 1 a,b. Average DNA content (a) and normalized AP activity (b) of control and experimenta l groups. Samples without BM P containe d considerably more DNA than controls (p0.2) at 10 days, but both SG-B P and SG-B I were significantly higher in normalized AP activity than the groups C, SG and C-BT (p 50%. Further reductions of such rate can only be achieved if the PE insert is replaced by a ceramic insert. The use of all-ceramic couples, i.e. THR femoral head and acetabular cup made of Biolox forte, allows for the rate of abrasion a reduction by factor 100 (tbl. 1) [6, 8]. The properties of Biolox (AI2O3) ceramics, which are described below, make the material particularly suited for hip endoprostheses: Aluminaceramicsare bioinert The high-purity alumina available today are extremely resistant to corrosion and will not emit any ions to the body. Hence, they can be classified as both, biocompatible and bioinert materials. This fact also means that they will not induce any connective osteogenesis. Wettability ofAluminaceramics The wettability of almnina in respect of polar liquids such as synovial fluid is better than the one offered by PE or CoCrMo. Hence, the material offers the best prerequisites for a lubricating film to establish between the prosthetic counterparts. Formation of the lubricating film is supported by the gap provided between the ceramic femoral head and the ceramic insert, which is generated by the difference in diameter between the head and the insert. The type of friction involved in vivo is a mixed friction which among other things is due to the oscillating movement of the hip joint. 165

166


Configuration fern, head/cup

Abrasion / year in mm

Metal/PE diam.

heads of 22 mm up to 0.5 mm rest: > 0.2 nun Biolox/PE 500 MPa. This means that the femoral heads and cup inserts made of it will notfracturewhen exposed to the different types of stress incurred in vivo. The resistance tofractureoffered by Biolox forte femoral heads of 28 nun in diameter exceeds 50 kN (> 5 tons), while the one offered by the smallest Biolox forte insert is 86kN(>8tons)[l,9]. However, the current experience in the use of ceramic cups has shown that the bone will not integrate the ceramic surface in the sense of a connective osteogenesis (ace. to Osbom), owing to the ceramic material’s absolute bioinertness. As a result, the implant is subject to early loosening. This basic property of ceramics is the cause of the failures experienced for monobloc cups. The autopsy examinations performed by Plenck [4], Hensge [3] and Fritsch [2] represent the connecting link between the theoretical and the in-vitro properties of all-ceramic couples, and the clinical results obtained for them. Such results can be summarized as follows: - Abrasion rates of < 0.005 nun/year are obtained if neither the stem nor the cup have loosened. - The diameters of the abrasion particles of AI2O3 ceramics ranges between 0.001 and 0.002 mm which is definitely lower than the ones observed for PE particles which were up to 0.02 mm. - The amount of abrasion particles is extremely low (< 0.005 mm/year) as long as there is no loosening of the implant and the cup and the stem. - Abrasion particles of AI2O3 ceramics are classified as biocompatible material. They will not cause any undesirable tissue reactions, which is due to their bioinertness and the small size and small amount of abrasion particles.

Improvementof THR With Spongiosa Metal Surface: G. Quack et al.

167

Osseointegratio n of the uncemente d SMS-TH R system The degree of osseous integration of uncemented metal cup components essentially depends on the design of the prosthetic surfaces. As far as the modular SMS cup system is concerned, the biological idea is to the fore: - Preservation of as much bone as possible. Resulting from this a spherical shape of tiie cup was devised. A cone is provided on tiie inside of the cup in order to allow for optimumfixationof the PE or of the ceramic insert. - Osseous integration of the prosthesis into the osseous bed through spongiosa-type wide-meshed enlargement of surfaces, which is achieved by coating the implant with tripodes: the cancellous bone represents the load-carrying and load-transmitting structure for positioning the cup and the stem. It reacts to the implant by integrating it. For this reason, the implant surfaces must feature a load-carrying structure and must provide suflBcient gaps into which the trabeculae of the spongiosa and the supplying vessels may vascularize [5]. Until December 1996, a total of 709 of these SMS cup systems were implanted. The follow-up examination of the first 95 uncemented hip endoprostheses (implanted in 1987 to 1991) with wide-meshed SMS cups yielded good medium-term results, with only one cup loosening and protruding into the minor pelvis (for data refer to table 2). The system used consisted of a ceramic femoral head (Biolox) and a PE cup insert. Yielded from this was a loosening rate of < 1.05% after a dwelling time of 6.03 years, and the good osseous integration predicted was confirmed. The consequences which can be drawnfromthefindingsobtained for the use of alumina ceramics [7] and from the positive medium-term results (*tained for the SMS cups must be a division of fiinctions [5]: - Ceramics should be used as gliding components offering minimum abrasion for the articular fimction as such. - Metal (CoCrMo, Ti) should be used for thefimctionof osseous integration. A cup system featuring a 3-dimensional wide-meshed surface and using Biolox forte ceramic inserts was developed, which as the central element offers a modular structure of the gliding couple. Resulting from this development was im imcemented and completely modular THRsystem offering optimum prerequisites to enable connective osteogenesis and low wear of the aJumina articular surfaces (fig. 1). On the basis of the experience gained from the use of the SMS cup system over the years, the improvement achieved in respect of the life of the prostheses is clearly due to the use of ceramic/ceramic couples. Also, it is possible to assign the causes of failures. Such findings shall be specified fiirther with the help of a comparative study investigating the use of BioloxÊ and Biolox forte/Biolox forte couples. SMS-CU P / aver. 6.03 y. - 95 SMS cups -47 females (49,47%) - 48 males (50,53%) Mean Score: Harris hip score 92.80 pts (+/- 7.11) Merle d’Aubigne sc. 16.74 pts (+/-1.36) - Aseptic loosening:! SMS cup = 1.05% TbL 2: Medium-term results

From November 1995 to March 1997, tiie first 50 patients below the age of 61 have received a gliding couple consisting of ceramic Biolox forte femoral head and insert. There have been no complications due to the use of ceramics. So far, the regular follow-up examinations have yielded good clinical (Harris hip score) and radiological results. In respect to the surgical technique used, the items specified below should be accounted for: - Use of an exact surgical technique and optimum positioning of the cup must be ensured.

168


- Both, the cone available on the inside of the cup, and the cone of the stem must never be damaged during handling and implantation . - Necessit y of using special implantation instruments . - To enable optimum cone-fitting , the ceramic insert must not be allowed to rest on the metal bottom of the cup (fig. 1). Summary The SM S THR-system (manufacture d by ESKA ) offers optimum osseointegratio n owing to the spongiosa-typ e design of its surface. The abrasion rate is drastically reduced as a result of the improved gliding propertie s of the all-ceramic (Biolox forte) couple, as long as the use of an exact surgical technique is ensured. The residual abrasion particles produced are bioinert and will be carried off by the organism without any problems. The ceramic component s will prolong the survival of the THR caused by the decrease of any osteolysi s due to abrasion. As a resuh the prosthesis will not loosen, or at least will loosen at a much later point in time. This means a much better prognosis, especiall y in the case of younger patients. A comparing study with the wear couples Biolox/PE and Biolox forte/Biolo x forte will help specifying these statements . First experience s with 50 cases, 2 years afler the first implantation , show good short-term results. Reference s 1. Clarke, I.C , P. Campbell, N. Kossovsky: Debris-mediate d osteolysi s - a cascade phenomeno n involving motion, wear, particulates , macrophage induction and bone lysis. In: St. John, K.R. (ed.): Particulate Debris from Medical Implants. AST M STP 1144, Philadelphia 1992, pp. 7-26. 2. Fritsch, E., H. Mittelmeier, J. Heisel, K. Remberger, S. Pahl: Micro- and macroscopic findings on capsular tissues of the hip after alumina arthroplasty. Proc. 6th Biomaterials Symposium "Ceramic Implant Materials in Orthopaedic Surgery", Sept 21-23, 1994, Gottingen (Germany); H.G. Buchhom, H.-G. Willert, in press 1996. 3. Hensge, E.J., I. Bos, G. Willmann: AI2O3 against AI2O3 combination in hip endoprostheses . Histologic investigation s with semiquantitativ e grading of revision and autopsy cases and abrasion measures. J. Materials Science Materials in Medicine 5 (1994) pp. 657-661 . 4. Plenck, jun. H., M . Buhler, A. Walter, K. Knahr, M. Salter: Fifteen years experienc e with alumina-cerami c total hip-joint endoprostheses : a clinical, historical and tribological analysis. In: Ravaglioli, A., A. Krjewski (eds): Bioceramic and the Human Body. Elsevier Appl. Sci., London, New York 1992, pp. 17-25. 5. (Juack, G., G. Willmann, H. Krahl, H. Grundei: Konzeptionell e Uberlegunge n zur Verbesserung der Pfanne der ESKA-Hftendoprothese durch die Gleitpaarung Keramik/Keramik [Conceptiona l consideration s relating to the improvemen t of the acetabula r cup of the ESK A hip endoprosthesi s achieve d through the use of ceramic/cerami c couples] . Biomed. Technik 41 (1996) 9, pp. 253-259 . 6. Saikko, V.: Wear test of the couple BIOLO X forte/BIOLOX forte. In: W. Puhl (ed.): Performance of the wear couple BIOLO X forte in hip arthroplasty. Enke Verlag Stuttgart, 1997. 7. Sedel, L., RS . Nizard, I. Kerbouli, J. Witvoet: Alumina-alumina hip replacemen t in patients yonger than 50 yars old. Clin. Orthop. 198 (1994), pp. 175-183 . 8. Walter, A.: Investigation of the wear couple BIOLO X forte/BIOLOX forte. In: W. Puhl (ed): Performance of the wear couple BIOLO X forte in hip arthroplasty. Enke Verlag Stuttgart, 1997. 9. Willmann, G.: Hiiftgelenkersat z - eine tribologische und konstruktive Herausforderun g [Hip arthroplasty - a challenge in respect of tribology and design]. Mat. Wiss. u. Werkstofitechni k 27 (1996) pp. 199-205 .


ACETABULA R ARTHROPLAST

RECONSTRUCTIO N IN REVISIO N TOTA L Y U S I N G A BON E G R A F T S U B S T I T U T E

HI P

R.P. Pitto and D. Hohmann

Department of Orthopaedics, Friedrich-Alexander University, Waldkrankenhaus St. Marien, RathsbergerstraBe 57, 91054 Erlangen, Germany ABSTRAC T Twenty acetabular reconstructions in revision total hip arthroplasty with severe loss of bone stock were performed combining the use of autogenous bone grafts, synthetical hydroxyapatite ceramic and reinforcement rings. The rings were fixed with screws on the host bone. The grafts fused within 4 months after the operation in all the cases. No migration of the acetabular component or lysis of the mixed graft was seen in 19 cases after 2 years. One implant failed because of malposition and was revised 6 months after the operation. These preliminary findings give rise to cautious optimism that this is a reliable method for acetabular reconstruction. KEYWORD S Hip Prosthesis, Revision, Bone Stock, Bone Graft Substitute, Ceramic. INTRODUCTIO N Deficiency of bone stock is a major problem in revision arthroplasty. Filling of the cavities by cement or metal leads to ftirther bone defects, if renewed loosening occurs. The use of autogenous bone grafts is a biological way to solve the problem, but the quantity of the available harvested material is limited. For ethical, bacteriological and viral safety reasons, management of bone banks is becoming increasingly restrictive [2]. Synthetical bone substitutes offer an alternative to homologous grafts. The goal of this study was to evaluate prospectively the clinical and radiological results of acetabular reconstructions after revision of loose acetabular components with severe bone stock defects combining the autogenous grafts with synthetical hydroxyapatite ceramic. MATERIAL S AN D METHOD S Twenty revision arthroplasties were performed using the impaction grafting technique on the acetabular side [4]. The morselized autogenous grafts were mixed with synthetical hydroxyapatite ceramic (granulate or blocks, Synthacerfi, Scientific Development, Munich, D) (Fig.l and 2). Reinforcement rings of Muller (9 cases), Ganz (5 cases) and Burch-Schneider(6 cases) (Protek, Munsingen, CH) were used to anchor the new Prothesis, to impact the mixed autoheterografts and to protect them during the healing. The clinical assessment was performed according to the criteria of the Chamley 6-6-6 Hip-Score-System [1] and a similar method was used to classify the pre- and post-operative bone stock of the acetabulum [3]: normal acetabulum (grade 6); peripheral ectasis (grade 5); protrusion (grade 4); ventral defect (grade 3); ventro-cranial defect (grade 2); dorsal defect or discontinuity of the pelvis (grade 1). 169


170

1 1

* ’’ ^ ^

*

^ 1 ^

ilk îi

% j’%^

Figure 1. The interconnecting porous framework of Synthacerfi, a synthetical hydroxyapatite ceramic [Ca5(Po4) 3OH].

Figure 2. The diameter of the pores of the synthetical hydroxyapatite ceramic is constant (600 jam). The porosity amount to 80%. Magnification: - = 100 |im.

Acetabular Reconstructionin Revision Total Hip Arthroplasty:R.P. Pitto and D. Hohmann 171

RESULTS The follow-up examination of the 20 patients at 24 months (min. 18, max. 28) showed the improvement of the pain score (av. pre-op, 2.6, av. Fw.-up grade 5.2), of the function score (av. pre-op. 2.5, av. Fw.-up grade 4.9) and of the motion score (av. pre-op. 4.3, av. Fw.-up grade 5.2). The roentgenological analysis of the grafts showed the fusion within 4 months after the operation in all the cases and a gradual condensation (Fig.3), but one of them had evidence of some degree of bony resorption. The bone stock had increased in all the cases (av. pre-op. 3.1, av. Fw.up grade 4.9). There were no signs of implant loosening. One case underwent a re-revision 6 months after the implantation because of malposition of the component and recurrent luxation. The graft showed in this case fusion at the host bone interface and osteointegration of the synthetical hydroxyapatite ceramic.

Figure 3. A) Aseptic loosening of a Wagner surface cup with severe osteolysis and protrusion. B) Revision and reconstruction of the acetabulum with impaction grafting and a reinforcement ring with hook of Ganz. C) Roentgenological signs of fusion of the graft, remodelling and stable implantation 2 years after surgery.

172


DISCUSSIO N The revision of the loose acetabular component with severe bony defect filled with a mix of synthetical hydroxyapatite ceramic and autologous bone grafts have proved to be of value. The results shows the good tolerance of the heterografts and a roentgenological evolution similar to that observed with pure autografts [5]. Further study is necessary, but these preliminary findings give rise to cautious optimism that this is a reliable method for acetabular revision, reconstruction and reconstitution. Careful pre-operative evaluation and peri-operative assessment to match bone defects, grafting patterns and reinforcement ring are of paramount importance. AKNOWLEDGMEN T The authors would like to express their thanks to Prof. K. Draenert and Dr. Y. Draenert, Centre of Orthopaedic Research, Munich, Germany, for the support of this study.

REFERENCES 1. 2. 3. 4. 5.

Charnley, J. In: Low FrictionArthroplastyof theHip, Springer, Heidelberg 1979, 20-24. Levai, J.P., Boisgard, S. Clin, Orthop.Rel, Res. 1996, 330, 108-114. Pitto, R.P. J. Bone Joint Surg.(Br.)(in print). Sloff, T.J.J.H., Huiskes, R., Van Horn, J. Acta Orthop.Scand. 1984, 55, 593-596. Stringa, G., Pitto, R.P., Di Muria, G.V., Marcucci, M. Int. Orthop. 1995, 19, 72-76.

POSTER 1



EFFEC T OF SOLUTIO N AGEIN G ON SOL-GE L HYDROXYAPATIT E COATING S B. Ben-Nissan, C.S. Chai and K.A. Gross Department of Materials Science, University of Technology, Sydney P.O. Box 123, Broadway, N.S.W., 2007, Australia

ABSTRAC T Sol gel technology offers an alternative technique for producing a bioactive surface for improved bone attachment. Hydroxyapatite was synthesized using the sol-gel technique with alkoxide precursors and the solution allowed to age up to 7 days. Coatings produced on MgO substrates were characterised by differential thermal analysis, thermal gravimetric analysis. X-ray diffraction and atomic force microscopy. It was found that, similar to the wet method of hydroxyapatite synthesis, an ageing time is required to produce a pure hydroxyapatite phase. KEYWORD S Hydroxyapatite, sol-gel, alkoxide, ageing, coating, characterisation INTRODUCTIO N Hydroxyapatite is an established material for applications such as maxillofacial reconstructive surgery and non load bearing applications [1]. One of the currently used methods to overcome low mechanical properties of bulk hydroxyapatite is to coat substrates such as titanium and its alloys. To date, many processes have been investigated. These include dip coating into a powder suspension [2], electrophoretic deposition [3], sputter coating [4] and plasma spraying [5]. Of these processes, plasma spraying is used commercially. Thermal spraying, however, requires good process control to avoid decomposition at high temperatures and is limited to coatings thicker than 30 jim. An alternative coating method is sol-gel deposition. While commonly being used for producing glasses and oxides, it has more recently been utilized to produce other more complex materials as well as non-oxide ceramics. The advantages of the sol-gel technique include (a) increased homogeneity due to mixing on the molecular scale, (b) reduced firing temperatures of ceramics due to small particles with high surface areas (c) ability to produce uniform fine-grained structures [6]. 175

176


Sol-gel techniques have been previously used to synthesize hydroxyapatite powder [7] and coatings [8,9]. This work illustrates the necessity of ageing time on the phase composition of the coating. METHOD S Solution Preparation 1.5x10"^ moles of calcium diethoxide (Kojundo Ltd., Japan) was suspended in ethanol and then dissolved in ethanediol (BDH Chemicals, Australia) with the aid of vigorous stirring in a glove box under dry nitrogen atmosphere. A second solution consisting of a stoichiometric amount of triethyl phosphite (Aldrich, U.S.A.) diluted in ethanol was prepared and added to the calcium bearing solution. Stirring was maintained for a period of ten minutes. Solutions were allowed to mature for 0 and 7 days before being used to make coatings. Coating Procedure Magnesia single crystal substrates (Zirmat, U.S.A.), 10x10x0.5 mm in size were chosen to study the coating quality, without the influence of the interactions with a reactive substrate such as titanium [10]. Substrates were ultrasonically cleaned in acetone and ethanol and then coated using a Headway Research (U.S.A.) spin coater. A volume of 0.5mL of solution was applied to the substrate and spun at 2500 r.p.m. for 10 seconds. Coated substrates were hydrolysed in an air oven (Labec, Aust.) at 70 C for 10 minutes, followed by prefiring at 500 C in a mufile furnace (Ceramic Engineering, Aust.) for 15 minutes. The coating/hydrolysis/prefiring procedure was repeated until 5 layers were deposited. After the final layer had been prefired, the coated substrates were heated at 200 C/hr to 1000 C and soaked for 15 minutes followed byfiunacecooling. Characterisation Technique s X-ray diffraction (Siemens D5000, Germany) was conducted on coated substrates using CuKa radiation and a glancing angle geometry. This attachment was necessary due to the small coating thickness. Scan parameters included a scan range of 28 to 40 29, step size of 0.02 , step time of 5 seconds and X-ray incident angles between 0.5 and 5.0 . Thermal analysis techniques differential thermal and thermogravimetric analysis (DTA and TGA) were performed using a SDT 2960 simultaneous thermal analyser (TA Instruments, USA). Samples were heated at 10 C/min to 500 C, held for 15 minutes and then heated to 1200 C at 200 C/hr (3.33 C/min). This heating rate was chosen to replicate the heating schedule. The morphology of the coated substrates were examined using a Park Scientific Instrument (Autoprobe LS, U.S.A.) atomic force microscope (AFM). RESULT S AND DISCUSSIO N The use of a pre-firing stage at 500 C facilitates coating build-up. It also removes the volatile species allowing the rapid heating rate to sintering conditions. Thermal shock is minimized due to the small coating thickness and the relatively small amount of material deposited. Hence, the thin coatings have a low susceptibility to thermal shock cracking, and facilitates ease of gas (including alchohol) removal. In addition, the thermal gradient within the coating is very small and the sintering conditions in all locations of the coating are similar.

Effect of Solution Ageing on Sol-Gel Hydroxyapatite Coatings: B. Ben-Nissan et al.

177

Fired coatings appeared quite uniform except towards the edges where it was thinner (seen as interference fringes). This would be a thinning of the coating due to the edge effect. Complete coverage is thus dependent upon the wettability and geometry of the object. Ageing The X-ray diffraction patterns for coatings produced after 0 and 7 day ageing periods are shown in Figure 1. Hydroxyapatite is evident after 0 days ageing, however, the presence of CaO (JCPDS 4-777) and other peaks suggests that the reaction has not reached completion. The coating produced after an ageing period of 7 days appears to consist solely of hydroxyapatite. Thus, it is evident that an ageing period is necessary to allow the different species present in the coating solution to mix thoroughly. Given the complex kinetics of this system, it is possible that some chemical reactions may take place during this maturing period. This ageing phenomena is similar to the ripening procedure used in the "wet method" to produce a stoichiometric hydroxyapatite [11]. Thermal analysis of the hydrolysed gel produced after maturing time of 7 days exhibited an endothermic peak at 110 C and three exothermic peaks at 216, 430 and 550 C. The large endotherm corresponds to the evolution of residual solvent and adsorbed moisture. This is followed by two large exothermic reactions at 216 and 430 C respectively. These reactions correspond to the formation of chemical bonds through condensation and polymerisation as well as the evolution of residual water and/or alchohol. This has also been reported with zirconia gels [12]. A smaller exothermic reaction occurs at 550 C. It is believed that this reaction represents the crystallisation of hydroxyapatite [13]. The vertical translation observed on the DTA/TGA curve at SOC’C represents the 15 minute pre-firing heat treatment. Surface Morphology The surfaces of the coatings were examined using AFM. The coatings were crack free and consisted of 2 distinct regions. The surface was covered with small grains, approximately 200nm in size. These smaller grains exhibited a "cauliflower-like" surface which was broken up by larger grains, approximately 800nm in diameter. These were observed at random separations across the coating surface and can be identified as peaks in figure 3a and lighter regions in figure 3b. It is possible that the larger grains had formed as a result of exaggerated grain growth. 1

’

1

’

I

’

i

1 S

^ ’c

I e 1

28

30

1

32

34

, 36

0 da>{s ageing 38

40

Degrees (20)

Figure 1. X-ray diffraction pattern of coatings produced from solution matured for 0 and 7 days.

200

400

600

800

1000

1200

Temperature (^C)

Figure 2. DTA/TGA plots for hydrolysed gel matured for 7 days.

178


A B Figure 3. Atomic Force Microscope scans of coatings, using solutions matured for 7 days. CONCLUSION S Hydroxyapatite coatings have been produced via the sol-gel route. It was found that to induce the formation of a coating that is predominantly hydroxyapatite, solutions should be aged prior to use. AFM examination revealed the presence of two distinct regimes consisting of grains 200nm and 800nm in size respectively after being sintered at 1000 C.

REFERENCES 1. 2. 3. 4. 5. 6. 7. 8. 9. 10. 11. 12. 13.

de Groot, K., de Putter, C, Sillevis Smitt, P.A.E. and Driessen, A.A. In : Scienceof Ceramics,Brit. Ceram. Soc, Stoke on Trent, 1981, 433-437. Lacefield, W.R., Ann.NY. Acad. Sci.,1988, 523, 72-80. Ducheyne, P., van Raemdonck, W., Heughebaert, J.C. and Heughebaert, M., Biomater., 1990, 11, 244-54. Ong, J.L., Lucas, L.C., Lacefield, W.R. and Rigney, E.D., Biomater.,1992,13, 249-254. Gross, K.A. and Bemdt, C.C., J. Biomed.Mat. Res.,to be published in 1997. Johnson, D.W. and Gallagher, P.K. In : CeramicProcessingbeforeFiring, John Wiley and Sons, U.S.A., 1978. Masuda, Y., Matubaram, K. and Sakka, S., J.Ceram. Soc. Japan,1990, 98, 1266- 1277. Chai, C, Ben-Nissan, B., Pyke, S. and Evans, L., In.SurfaceModificationTechnologiesVII, T.S. Suddshan, K. Ishizaki, M. Takata and K. Kamata, Eds. Cambridge University Press, UK, pp. 509-525, 1994. Deptula, A., Lada, W., Olczac, T., LeGeros R.Z. and LeGeros J.P., In : BioceramicsVol. 9 University Press, Great Britain, 1996, 313-316. Chai, C. and Ben-Nissan, B., J. Aust. Ceram. Soc, 1993, 29 (1/2), 81-90. Osaka, A., Miura, Y., Takeuchi, K., Asada, M. And Takahashi, K., J. Mat. Sci.:Mat. in Med.,1991,2,51-55. Ben-Nissan, B., Anast, M., Bell, J., Johnston, G., West, B.O., Spiccia, L., de Villiers, D. and Watkins, I., Proc. 1stInt. Symp. Sci. ofEng.Cer.,S. Kimura and K. Niihara Eds., MikawaHaitsu,Koda, 1991,25-29. Gross, K.A., The amorphous phase in hydroxyapatite coatings, PhD dissertation, 1995, State University of New York at Stony Brook, USA.


IONI C CEMENTS : INFLUENC E OF LIQUID/SOLI D RATI O ON POROSIT Y AND MECHANICA L PROPERTIE S F.Betchem*, P. Michaud*, F. Rodriguez*, Z. Hatim**. * Laboratoire de Pharmacie Galenique, Faculte des Sciences Pharmaceutiques, 35, chemin des Maraichers 31062 Toulouse Cedex ** Laboratoire de Chimie-Physique, Universite Chouaib Doukkali, Faculte des Sciences, B.P. 20 El Jadita-Maroc. ABSTRAC T Ionic cements are widely studied in orthopaedics and the biocompatibility, nontoxicity, and partial resorbability of hydroxyapatite are well known. In this study, we analyzed the effect of varying the liquid/solid ratio over a short range (0.40 - 0.50) with an ionic cement which has Ca/P atomic ratio 1.63. We examined the influence on axial and diametral tensile strengths, hardening and porosity. We observed that, when the amount of liquid is increased, the excess water is not used for the reaction, and occupies an interstitial position in the solid. So the porosity and the hardening time increase while the tensile strength decreases. KEYWORD S Ionic cements, hydroxyapatite, porosity, hardening, mechanical resistance. INTRODUCTIO N Ionic cements are being increasingly studied in orthopaedics. Their interest is due to their composition and structure, which are close to those hydroxyapatite (Caio(P04)6(OH)2, HAp), their easy utilization, and their non-exothermicity [1,2]. In order to reach total rehabitation, a cement has to present sufficient porosity to allow diffusion of body fluids but an increase in porosity is frequently correlated with a decrease in mechanical resistance. This paper shows how the variation of the liquid/solid ratio over a short range (0.40 - 0.50) leads to modifications of the mechanical and physical properties in vitroand in a moist atmosphere at 37 C. The value Ca/P of 1.63 was chosen in order to prepare a non stoechiometric apatite and to be sure to obtain a total reaction of all the initial reagents. MATERIAL S AND METHOD S Cement paste was prepared by addition of a liquid containing calcium and phosphate ions to a mixture of solid Calcium Phosphates (table 1) Ca4(P04)20, TTCP and a-Ca3(P04)2, a-TCP [3]. In the cement formula, the Sodium Glycerophosphate (NaGP) is useful for improving the paste homogeneity. The paste was placed in silicone molds (9 mm diameter and 5 mm thickness) for 10 minutes in a moist environment at 37 C. Then, the cylindrical samples were removed from the molds and kept in a moist environment at 37 C for 96 hours. Before analysis, each sample was dried in an infrared balance at 110 C for 10 minutes. These samples were used to determine the mechanical properties with a Diametral Tensile Strength (DTS) machine PHARMATEST PTB 311 and computerized single punch machine KORSH Ekod for Axial Tensile Strength (ATS). The specific area and total pore volume of the samples were measured by a nitrogen high speed surface area 179

180


Powder

12000 7

8.68 g 1.32 g

TTC P and a-TCP NaGP

E E 10000 8000

Liquid

Ca(OH)2 H3PO4 H2O

Ca/P=1.635 liq/sol=0.4 3 "Ca/P=1.635 liq/sol=0.4 5 - Ca/P=1.635 Iiq/sol=0.5 0 B

6000 +

0.12g 0.29 ml to 100%

.5 4000 % 2000 24 30 36 Time (min)

Table 1: Composition of the cement

Graph 1 : Hardness versus time (Ca/P = 1.63, hquid/solid = 0.43) and pore size analyzer Quantachrome NOVA 1000 for specific area; mercury pore size analyzer Micromeretic autopore 11-9215 for total pore volume. To follow the hardening, we used a texture analyzer TAXT2. Cement paste was introduced into a mold of 12 mm diameter and 12 mm height which was placed in a steel block of great thermal inertia to keep the cement temperature at 37 C during measurements. One measurement was made every 3 minutes for one hour with a 1 mm punch in the following conditions : 2 mm/s for downward and rise speed; penetration of 5 mm. For each cycle we took the maximum strain to plot the hardening curve.

RESULT S AND DISCUSSIO N The graph 1 shows the setting and the hardening evolution of a cement with three different liquid/solid ratios in a moist atmosphere at 37 C. In the three curves, there is two main stages before reaching to hydroxyapatite. The part A corresponds to the formation of brushite (CaHP04, 2H2O, DCPD) by acidic and basic reactions between TTCP, a-TCP and phosphoric acid contained in the solution. 120 J

ir no f B

S 100 ^

^

^ g.

51 *Z 49 ^

70 f 60

0.35 0.43 045 Uquid/solid ratio

93.37 53’ 83.53 49

90 +

59 111.94 + 57 ’ 56 101.26 55 ^ 54

47 H-

-H

0.4

0.45

-+0.5

45 0.55

Liquid / solid ratio

Graph 2 : Tensile strength after 96 hours in Graph 3 : Porosity and specific area versus moist atmosphere at 37 C liquid/solid ratio Liquid/Solid ratio 0.40 0.43 0.45 0.50

Apparent density 1.56 1.50 1.48 1.38

Measured density 2.46 2.46 2.52 2.49

Theoretica l % of open porosity 37 40 42 45

Residual humidity rate (% ) 12.50 14 15 15.50

Table 2: Theoretical open porosity and residual humidity rate of sample differents from their liquid/solid ratio.

Ionic Cements:Influenceof theLiquid/SolidRatio:F. Betchemet al. 181 00

0.3

’JS^:

0.2 2 0.1

S 0.0-H^ ^ 00

o

H-3

rrr rr 10000

1000

100

Diameter (Angstroms)

s (view by Figure 1: Distribution of pore size by mercury Figure 2 Porosity and HA p needle intrusion in 0.43 liquid/solid ratio sample. SEM X 20000) in 0.43 liquid/solid ratio sample. The part B correspond s to the formation of octocalciu m phosphate (Ca8H2(P04)6, 5H2O, OCP ) [4] which in turn gives HA p after several hours. The curves also show that the setting and hardening times increase as the liquid/solid ratio increases . In graph 2, we observe the large modificatio n of mechanica l propertie s when the amount of liquid is increased . Graph 3 shows the rise of specific area and porosity as the liquid/solid ratio increases . The rise of porosity is due to excess water which occupies an interstitial position during hardening in a moist environmen t at 37 C. When the liquid is removed by drying in an infra red balance it contribute s to the porosity and specific area. Table 2 shows the results of density and humidity rate measurements . The differenc e betwee n apparent and measured density is explaine d by the open porosity. Moreover the rise of the residual humidity rate as the liquid/solid ratio increase s shows that the reaction which leads to HA p uses the necessar y amount of water and the excess of water occupies an interstitial position in the solid. The measuremen t of excess of liquid consists in measuring the residual humidity rate. The results achieve d by nitrogen and mercury pore size analyzers are compatible with the theory: in each sample the rate of open porosity depends on liquid/solid ratio and is included betwee n 45 and 60%. The decrease of mechanica l resistance is correlate d with porosity [5,6]. Another fact can contribute to the decrease of the mechanica l propertie s : a large amount of liquid slows down the crystallizatio n by modifying the different calcium phosphate s successivel y formed. The mercury pore size analyze of samples shows that there is two pore families for each liquid/solid ratio : the main family is situated near 1000 Angstroms diamete r (figure 1) and increase s as the amount of liquid increases . The vision of figure 2is a SEM photograph (x 20000) illustrating the porosity and the needle s of HA p inside the cement (Ca/P =1.63 and liquid/solid ratio = 0.43). CONCLUSIO N The liquid/solid ratio is an important paramete r in the preparation of ionic cements . The amount of water required is determine d by the Ca/P ratio. Excess liquid is not used for the reaction but occupies an interstitia l position in the solid. When the excess of liquid increases , the porosity and the hardening time increase and the tensile strength decreases . To achieve sufficient porosity while maintaining correct mechanica l properties , the formulation of those cement s must be optimize d and the precise conditions of their utilization must be determined .

182


ACKNOWLEDGMEN T We thank TEKNIMED (B.P. 60, Vic en Bigorre, France) for financial support of the work and for providing of powder raw materials. One of the authors (F.B.) wishes to gratefully acknowledge Pr J.L. Lacout for encouragement and support.

REFERENCES

1. E.W. BROWN, L.C. CHOW, J. Dent Res.,62, 1983, 672. 2. A. A MIRTCHI, J. LEMAITRE, E. MUNTING, Biomaterials,11, 1990, 83-88. 3. J.L. LACOUT, E. MEJDOUBI. Procede d’obtention d’hydroxyapatite phosphocalcique, application au comblement osseux ou au moulage de pieces et produits utilises. Brevet Fr92.09019/PCT/FR. 4. J.L. LACOUT, E. MEJDOUBI, M. HAMAD, J. Mater Sci, Mater.Med, 7, 1996, 371374. 5. R.WmCE, J. Am. Ceram. Soc, 79, 1993, 1801-1808. 6. O. BERMUDEZ, M.G. BOLTONG, F.C.M. DRIESSENS, J.A. PLANELL, J. Mater. Sci. Mater.Med, 4, 1993, 389-393.

Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd

SINTERIN G BIOCERAMIC

AND S

THERMA L

DECOMPOSITIO N

OF

HYDROX Y APATIT E

J. Cihlar and M. Trunec Department of Ceramics, Institute of Materials Engineering, Brno Technical University, Technicka 2, 616 69 Brno, Czech Republic

ABSTRAC T In the course of high temperature treatment of injection moulded hydroxyapatite ceramics (HA) sintering, grain growth and thermal decomposition of HA to tricalcium phosphate took place. The sintering was finished at 1573 K. The grain growth started at 1500 K and the thermal decomposition started at 1623 K. The activation energy of grain growth was 215" kJ/mol, that of thermal decomposition 283.5 kJ/mol. The optimum sintering temperature was found at 1473 K. KEYWORD S Hydroxyapatite ceramics, thermal decomposition, gram growth, sintering, kinetics INTRODUCTIO N Properties of sintered hydroxyapatite ceramics, namely their mechanical and biochemical properties depend on the physical and chemical structure of HA [1]. This structure is dependent on processing parameters of HA ceramics, namely on conditions of thermal treatment. In the course of thermal treatment sintering, grain growth and thermal decomposition of HA take place [2]. Data published about the physical and chemical behaviour of HA ceramics in the course of thermal treatment are not consistent. Nonuniformity namely has to do with an optimal sintering temperature and mechanism of thermal decomposition of HA ceramics [3, 4, 5]. In this contribution the authors try to make a kinetics and mechanism of thermal decomposition more clear and to give optimal tempei-ature of HA sintering. MATERIAL S AND METHOD S The samples of HA ceramics were prepared by ceramic injection moulding. Materials used and sample processing had been published [2]. Thermal treatment of HA samples was made in a superkanthal furnace in air atmosphere of 25% relative humidity at the temperature range from 1373 K to 1773 K for 1 to 22 hours. The microstructure and microanalysis of HA specimens were determined by SEM on a JXA-840 microscope equipped with an energy dispersion analysator (Link). The phase composition of HA ceramics was established by X-ray diffraction analysis on a D-500 diffractometer (Siemens). For quantitative phase analysis the part of spectrum in the range of 30 to 50 for 26 was used. The diffraction of (121) + (211), (300) and (301) planes were used for HA content determination, the diffraction of (123) + (254), (434) + (264) + (401) and (400) planes were used for the determination of tricalcium phosphate (TCP) content. 183

184


RESULTS AND DISCUSSION Sintering The dependence of linear shrinkage of hydroxyapatite ceramics on sintering temperature is given in Figure 1. In the temperature range from 1373 to 1573 K linear shrinkage increased with increasing temperature. A maximum shrinkage (16%) was obtained at the temperature of 1573 K. With this maximum linear shrinkage the hydroxyapatite ceramics had relative density of 98%. Sintering at the temperature above 1573 K (in the air atmosphere) did not result in further increase of the density of hydroxyapatite ceramics.

1100

1200

1300

1400

SINTERING TEMPERATURE [K]

Figure 1 The dependance of the shrinkage of HA ceramics on sintering temperature

Figure 2a Microstructure of injection moulded HA ceramics sintered at 1373 K for 1.5 hour SINTERING TIME [hour] ’ \m Y

2

4

1

1

’

6

1

-I

8

T

1

Sintering Time Sintering Temperatur e

1

\

iy

1

10

A

Y-

0 1000

* t-^t

1200

1

1

-1 5.

1400

\

1600

SINTERING TEMPERATURE [K]

Figure 2b Microstructure of injection moulded HA ceramics sintered at 1773 K for 1.5 hour

Figure 3 The dependence of the grain size of HA ceramics on sintering time and temperature

Sintering and ThermalDecompositionof Hydroxyapatite Bioceramics:J. Cihlar and M. Trunec 185

Grain Growth The microstructure of the sintered HA ceramics is shown in Figure 2. The average grain size of the HA ceramics sintered at 1373 K for 1.5 h was about 1 |Lim (figure d). The HA ceramics sintered at 1773 K for 1.5 h contained grains of the average size of 16 fin (^Figure 2b). The growth of grain size depended, above all, on sintering temperature. The most pronounced grain growth was observed in the temperature range from 1573 K to 1673 K (see Figure 3). Thermal decomposition The loss in weight (due to the loss of water) of hydroxyapatite ceramics started ai the temperature of 1373 K. A negligible shift of diffraction lines of HA iccompanied b> weight loss of HA was caused by formation of oxyapatite (OA). HA-OA system (termed as hydroxyoxyapatite [4]) was stabile for 15 h at temperature 1573 K. The presence of crystalline a-TCP was detected until sintering for 2 hours at 1623 K. The course of thermal decomposition is perceptible from Figure 4.

Figure 4 Section by the layer of TCP on the surface of HA ceramics sintered at a) 1573 K for 22 hour b) 1773 K for 8 hour %

200

1 1 1 1 1 1 1

1

LU

150 Q. O

^

CO CO

UJ

z "^ o X

1

1

1

1

\

-4 ^ J

c

\

A 1

1

^

/ Oi [ 0

1

-5

yi

50

1

Ep = 283. 5 kJ/mo l

^^^^"^^ 1

100

O

r - - i-

I ’ 1 ’ 1 1

1

1 1

I

.

I

.

I

.

200 400 600 800 100012001400 SINTERING TIME [min]

Figure 5 The dependence of the thickness of TCP layer on the sintering time

1

0,56

0,57

.

1

0.58

1

I

0,59

.

I

0,60

.

0,61

lOOOyT [1/K] Figure 6 The temperature dependence of the rate constant of HA decomposition

186


The thermal decomposition started on the surface of HA ceramics. At first, islands of TCP appeared (Figure 4a). These islands were connected together in the compact TCP layer growing into the inside of HA-OA ceramics (Figure 4b). The growth of thickness of TCP layer (x) with time (t) was described by parabolic rate law x=(Kt)^^^ (see Figure 5) [6], where K is a factor of proportionality. The rate of the thermal decompositicm of HA-OA is then controlled by diffusion of reaction products (water) through the layer of TCP. The activation energy of thermal decomposition of HA-OA was 283,5 kJ/mol (see Figure 6). In the temperature range from 1623 to 1773 K, thermal decomposition products of HA ceramics contained only TCP. Traces of tetracalciumphosphate appeared at 1773 K. The thermal decomposition of HA ceramics could be described by equations: 2 Ca5(P04)30H o Caio(P04)60 + H ^ (1373 - 1723K)

(1)

Ca,o(P04)60

HA

HA

Fig. 2b Higher power magnification of the rectangular area seen in Fig. 2a. There is no intervening fibrous tissue between bone (B) and HA surface. Osteoblastic cells (arrows) are seen on the HA surface.

236


inductive action. Because the mesenchymal stem cells reside in periosteum, the stem cells migrated from the periosteum into the pore regions can show new bone formation on the HA surface. It is also well known that the existence of BMPs in bone tissue (bone matrix), and as described above, exogenously added BMPs can induce osteoblastic differentiation primarily on the HA surface. Based on these results, we suppose mesenchymal stem cells from surrounding periosteum and protein factors (BMPs) in newly formed bone tissue can participate in the osteogenesis in HA pore areas. When bioactive materials such as HA was implanted in bony defects, Osbom reported[ 12] that bone bonding was accomplished through the cascade of bonding osteogenesis. That is, the new bone formation occurs on the bioactive material’s surface. Though the mechanism of the bonding osteogenesis is not ftiUy understood, the preexistence of bone tissue near the implantation sites is the source of osteogenic cells or osteoinductive factors to show new bone formation. The phenomenon of the new bone formation is also known as osteoconduction. The term of osteoconduction implies the appearance of new bone tissue around implanted materials. When the implanted materials are bioactive, as reported by Osbom, the bone bonding occurs by the cascade of bonding osteogenesis and when the materials are bioinert, the bone contact occurs by the cascade of contact osteogenesis [12]. Therefore, osteoconduction is the phenomena at orthotopic sites which leads to new bone formation around implanted materials not related to the materials properties. The present experimental model of periosteum/HA composite which was implanted at subcutaneous sites also showed new bone formation around the implanted materials. The bone was derived from the periosteum which surrounded the HA and therefore, present subcutaneous implantation of periosteum/HA can be regarded as an osteoconduction model at ectopic site. Importantly, the bone formation in the pore regions initially occurs on the HA surface (bonding osteogenesis) and suggest that present osteoconduction model can be available to evaluate the materials properties regarding bone bonding. REFERENCES 1. Ohgushi, H.,Okumura, M.,Tamai S.and Shors,E.C. J. Biomed.Mat.7^^^. 1990,24,1563-1570 2. Okumura, M., Ohgushi H. and Tamai, S,Biomatenals1991,12, 411-416 3. Ohgushi, H., Okumura,M., Yoshikawa, T., Inoue, K., Senpuku, N. and Tamai, S., J. Biomed Mat. /?e5.1992,26,885-896 4. Yoshikawa, T., Ohgushi, H. Okumura, M., Tamai, S., Dohi, Y. and Moriyama, T., Calcified TissueInternational1992 50,184-188. 5. Ohgushi, H.,Dohi, Y.,Tamai, S.and Tabata, SJ.BiomedMat. Res.1993.27,1401-1407. 6. Ohgushi, H., Okumura, M. Yoshikawa, T. Tamai, S. Tabata, S and Dohi, Y. in BonebondingBiomaterials,Helthcare Comm. Publ, the Netherland, 1992,pp.47-56. 7. Takaoka, T., Okumura, M., Ohgushi, H., Inoue, K., Takakura Y. and Tamai, S. Biomaterials1996,17,1499-1505 8. Ishida, H, Tamai, S., Yajima, H., Inoue, K., Ohgushi, H. and Dohi, Y., Plast. Reconstr. Surg. 1996,97,512-518 9. Ishida, H., Ohgushi, H., Inoue, K., Yoshikawa T., Yajima, H., Tamai, S. and Y. Dohi. Bioceramics 1996,9,73-76 10. Ohgushi, H., Okumura, M., Inoue, K., Dohi, Y.,Tamai, S.Murata, M., and Kuboki,Y. Bioceramics 1995,8,61-67 11. Caplan, A.I. Clin. Orthop. 1990,261,257-267 12. Osbom, J.F. and Newesely, H. Biomaterials 1980, (John Wiley and Sons, Ltd, edts: G.D.Wmter, D.F. Gibbons and H. Plenk Jr.) 1982, 51-58

Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd

CONDITION S OF THE COPRECIPITATIO N OF CALCIU M HYDROXYAPATIT WIT H ZrOi , ZrOi+YjOa , AI2O3 FROM AQUEOU S SOLUTION S USIN G AMMONI A

E

V.P. Orlovskii, Zh. A. Ezhova and E.M. Koval Institute of General and Inorganic Chemistry of Russian academy of sciences, Moscow, Leninskii pr. 31, 117907 Russia

ABSTRAC T The systems CaCl2-Zr(OH)2Cl2-(NH4)2HP04-NH3-H20, CaCl2-AlCl3-(NH4)2HP04-NH3-H20, CaCl2-YCl3-Zr(OH)2Cl2-(NH4)2HP04-NH3-H20 are studied at the 25 C using the solubility method (Tananaev method of residual concentrations). The conditions of coprecipitation of calcium hydroxyapatite (HA) with hydroxides of metals (Zr; Al; Zr with addition of Y) are determined. Using the chemical analysis, XRDA and IRS methods it is shown that after the coprecipitated phases calcined at 900 C the homogenous mixture of HA and oxides (Zr02; AI2O3; Zr02+Y203) is formed. Zirconium dioxide is crystallized in tetragonal modification. KEYWORDS : hydroxyapatite, zirconium dioxide, aluminium oxide, coprecipitation INTRODUCTIO N The preparation and detailed physical-chemical investigations of new phases based on HA of various dispersity degree and AI2O3, Zr02, Zr02+Y203 and other is one of fundamental problem of HA chemistry and technology. The addition of appropriate additives improves the mechanical characteristics of HA bioceramics without decreasing their biocompatibility. Previously, it was used the solubility method to study interaction in the CaCl2-(NH4)2HP04-NH3-H20 system at 25 C and determined the optimum conditions for obtaining HA. The major factors involved in securing pure HA (free of tricalcium phosphate TCP and other calcium phosphates) are the ratio between the initial components (Ca/P), pH, the time of attainment of equilibrium, and the order of mixing solutions [1]. A high value (^10) favors the coprecipitation of HA and zirconium and yttrium hydroxides [1,2]. In present work interaction is studied by the solubility method (residual concentrations) in following systems at 25 C: CaCl2-Zr(OH)2Cl2-(NH4)2HP04-NH3-H20; CaCl2-YCl3-Zr(OH)2Cl2-(NH4)2HP04-NH3-H20; CaCl2-AlCl3-(NH4)2HP04-NH3-H20. 237

238


MATERIAL S AND METHOD S Aqueous solutions of CaCb, AICI3, YCI3, Zr(OH)2Cl2-8H20 (twice recrystallized from water-alcohol solutions of hydrochloric acid), (NH4)2HP04, and ammonia were used as initial components. The solutions were prepared with the use of boiled twice-distilled water. Interaction in the system was studied at a constant ammonium phosphate concentration (0.025 mol/1), and variable amounts of calcium, yttrium, aluminium and zirconium ions. Ratios in the initial mixture were: ni=CaCl2/(NH4)2HP04=1.67^1.75 n2=Zr(OH)2Cl2[(Zr(OH)2Cl2+2YCl3);2AlCl3]/(NH4)2HPO4=0.033^.2 n3=Zr(OH)2Cl2/(Zr(OH)2Cl2+2YCl3)=0.85 and 0.97 n4=2YCl3/(Zr(OH)2Cl2+2YCl3)=0.15 and 0.03 n5=NH4OH/AlCl3=3.0-J-4.0. Systems with Zr(OH)2’^-ions was investigated at pH-lO, which attained by adding to each specimen a concentrated aqueous solution of ammonia. The total volume of the mixtures was 200 ml. The mixtures were vigorously stirred at 25–0.1 C until equilibrium was attained. Then solutions were filtered, the pH was measured and the chemical analysis of the liquid and solid phases were carried out. RESULT S AND DISCUSSIO N I. Systems CaCl2-Zr02Cl2-(NH4)2HP04-NH3-H20 and CaCl2-YCl3-Zr(OH)2Cl211(NH4)2HP04-NH3-H20. To the solution containing CaCl2 and concentrated ammonia were added at first the (NH4)2HP04 and then the Zr(OH)2Cl2 (Zr(OH)2Cl2+YCl3) solutions. In this case the precipitate formed immediately. The equilibrium was reached for 7 days. After the stirring of mixture for one day in liquid phase were detected the considerable amounts of calcium ions; the phosphate and zirconium (zirconium and yttrium) ions practically absent. Further stirring of mixtures for 7 days leads to a gradual quantitative transition of calcium ions into the solid phase. The Ca/P ratio in the solid phase nsoiid=1.67 (pH=9.8) indicates the formation of mixed calcium hydroxyapatite and zirconium (zirconium and yttrium) hydroxide(s) phases. The composition of resulting solid phases is described by the general formulas: Caio(P04)6(OH)2-Zr(OH)4-xH20 Caio(P04)6(OH)2-m{ [Zr(OH)4]o 97[2Y(OH)3]o 03}xH2O Caio(P04)6(OH)2.m{[Zr(OH)4]o85[2Y(OH)3]o,5}-xH20, where m=0.2-^1.2; x=6-12. Heating of hydrated mixed phases of HA and zirconium (yttrium) hydroxide to 900 C, as shown the IRS and XRDA data, leads to the gradual water removing. HA not decomposed and not interacted with the formed zirconium dioxide and yttrium oxide. The HA structure is conserved. Homogeneous mixtures of HA and zirconium dioxide, HA, zirconium dioxide and yttrium oxide with mixing composition formed after calcination at 900 C: Caio(P04)6(OH)2-mZr02 Caio(P04)6(OH)2-m[(Zr02)n(Y203)i-n], where m=0.2^1.2; n=0.85; 0.97.

Coprecipitationof HA With Zr02, Zr02 + Y2O3, Al20sfrom Aquoeus Solutions: V.P. Orlovskii ti al.

4000

2000 1000 Wavenumber (cm’^) Figure. XR D pattern (a) and IR spectrum (b) of 3Ca3(PO)4Al203. Marks : the lines due to AI2O3 (solid sircle)

10

20

30

40

50

60

239

400

X-ray powder diffraction data showed that in all calcinate d solid phases of HA (without the TC P and CaO impurities) and Zr02 in the tetragonal syngony presented . The IR absorption spectra analysis shows that in obtained phases the PO^-ions are strongly distorted, as is indicated by a significant frequenc y split. The spectra show a clearly defined, narrow band of stretching vibrations of the Oir"-groups v(OH>-3570 cm’\ After the heating at 900C the water bending vibrations 5(H2O)=1650 cm"^ disappear, whereas the absorption bands of PO J and OH" remain practically unchanged. The IR absorption spectra of dehydrate d phases are analogous to the IR spectra of HA [3]. II . System CaCl2-AlCl3-(NH4)2HP04-NH3-H20 (25 C). The amphoteric propertie s of Al were considere d in the investigatio n of the system with AICI3 . Ammonia was added in the system in summary amounts neede d for HA formation of the reaction: 10CaCl2+6(NH4)2HPO4+8NH4OH -> iCaio(P04)6(OH)2+20NH4Cl+6H2 0 (CjgH3react~^-^^32 5 mol/1

-

coust)

and

in

agreemen t

with

the

variable

ratio

n5=OH/Al=3.0-^4.0 . With the exceptio n of AIPO 4 formation, alumunium chloride was added in end of precipitation . Interaction is system depends on pH precipitatio n and goes in several studies. 1. At the points of the system with nl=1.67-^1.75 , n2=0.166 , n5=3.0-^3. 2 and pH=7.64-8. 0 at the stirring of initial component s for 14 days (and more) in liquid phase was detecte d a significant quantity of calcium ions; aluminium and phosphate ions practically absent. The Ca/P ratio in the solid phase correspond s to nsoiid=1.5 . In IR absorption spectra of isolated in this region solid phases the stretchin g vibrations of OH~-group v(OH) at 3570 cm"^ is absent. Therefore HA on this precipitatio n study not formed and reaction betwee n CaC b and (NH4)2HP04 proceede d with the TCP formation. Precipitate d hydrated phase consiste d of TCP and A1(0H)3 with the small amount of CI" - ions after the calcination. If calcined at 900 C, this phase formed the mixed phase that consist of TCP and AI2O 3 of the compositio n 3Ca3(P04)2-Al203 . The X-ray powder diffraction data (Fig.a) and IR-spectrum of 3Ca3(P04)2-Al20 3 (Fig.b) are analogous to that of >S-Ca3(P04)2 synthesize d by as [4].

240


On the diffraction pattern all peaks characteristics for >S-Ca3(P04)2 are present. Absence of band v(OH) at 3570 cm"^ and splitting of stretching (V3) and bending (V4) vibrations of P04’ tetragon are characteristic to IR-spectra of this phase. The band of P-0 antisymmetric valence vibrations V3 is splitted on some components with a maximum 1116, 1090, 1080, 1038, 1020 cm\ The band of vj vibrations are 969 and 940 cm"\ The band of bending vibrations 0-P-O are V4 - 600, 590, 549, 540 cm’\ 2. At the increasing of precipitation’s pH 8.09.55 (n5=40) and stirring of initial components for 14 days calcium and phosphate ion were not detected in liquid phase, however aluminium presented in significant amount. Ratio nsoiid=Ca/P=1.67 therefore in solid phase HA presented without the TCP impurity. Aluminium precipitated in form of variable composition’s basic salts. For example at ni=1.67;n2=0.083-H0.2; n5=4.0; pH=9.55 the precipitates of composition Caio(P04)6(OH)2-2[(NH4)x-yAl(OH)3y+x]-zH20, y=0.3-i-1.0 formed in solid phase. If calcined at 900 C,the solid phases had the composition Caio(P04)6(OH)2-yAl203, where y=0.3^1.0. CONCLUSION S Interactions in the CaCl2-Zr(OH)2Cl-(NH4)2HP04-NH3-H20, CaCl2-YCl3-Zr(OH)2Cl2(NH4)2HP04-NH3-H20 systems at 25 C are studied by the method of residual concentrations variant. The investigation of these systems [5, 6] are given the wide information about the optimal conditions of HA and covalent metals hydroxides coprecipitation, the compositions of precipitated phases and the calcination products. The results of these investigations may be used for the development of bioceramics with assigned and determined mechanical properties. REFERENCE S 1. Orlovskii, V.P., Ezhova, Zh.A., Rodicheva, G.V., Koval’,E.M., Sukhanova, G.E., and Tezikova, L.A., Zh.Neorg, Khim.1992, 37, 4, 881. 2. Ezhova, Zh.A., Rodicheva, G.V., Koval’,E.M., and Orlovskii, V.P., Zh,Neorg. Khim. 1991, 36, 10, 2494. 3. Chumaevskii, H.A., Orlovskii, V.P., Ezhova, Zh.A., Minaeva, N.A., Rodicheva, G.V., Steblevskii, A.V., and Sukhanova, G.E., Zh. Neorg. Khim.1992, 37, 6, 1455. 4. Ezhova, Zh.A., Orlovskii, V.P., and Koval’,E.M.,., Zh. Neorg. Khim.1997, 42 (in press) 5. Ezhova, Zh.A., Orlovskii, V.P., Koval’,E.M., and Kozhenkova, E.B., Zh.Neorg. Khim.1996, 41, 11, 1686. 6. Ezhova, Zh.A., Orlovskii, V.P., and Koval’,E.M.,., Zh. Neorg. Khim.1995, 40, 10,1563.


TRANSFORMATION OF a -TCP TO HYDROXYAPATITE IN ORGANIC MEDIA Kiyoko Sakamoto^, Shunro Yamaguchi^, Atsushi Nakahira^, Kazunori Kijima^, and Masayuki Okazaki"^ ^ Department of Chemistry, Osaka Sangyo University, Nakagaito, Daito, Osaka 574, Japan, ^ ISIR, Osaka University, Mihogaoka, Ibaraki, Osaka 567, Japan, ^ Kyoto Institute of Technology, Goshokaidocho, Matsugasaki, Sakyo-ku, Kyoto 606, Japan, "* Osaka University Faculty of Dentistry, Yamadaoka, Suita, Osaka 565, Japan.

ABSTRACT The transformation of ot-tricalcium phosphate (ot -TCP ) to hydroxyapatite ( HAp ) in organic media has been investigated. Hydrolyses of ot-TCP in a series of aliphatic alcohols were carried out under the control of pH and the reaction temperature. The transformation rates and microstructures of HAp were influenced by the hydrophobicity of the aliphatic alcohols. The formation rates of HAp increased with increasing hydrophobicity of alcohols. The rates in 1octanol and 1-hexanol were 5-times faster than that in the hydrophilic alcohols and were compatible to that in the absence of alcohols. The microstructures of HAp prepared in the hydrophobic alcohols were the needle-like particle (length ; l.O’^Z.O U m) and differed from HAp prepared in hydrophilic alcohols and in the absence of alcohol. KEYWORD S ; OL -tricalcium phosphate, hydroxyapatite, transformation, aliphatic alcohols. INTRODUCTIO N There has been considerable interest in transformation of ot-tricalcium phosphates ( a -TCP ) to hydroxyapatite ( HAp ) due to the convenient control of crystal formation. The formation of HAp in the hydration and hardening of a-TCP has been extensively investigated by H. Monma et al. [1,2,3]. While, in organism the formation process of HAp is extremely complication due to concern with organic compounds. It is known that other inorganic phosphates were effectively prepared in organic media [4]. Therefore, the transformation process of OL -TCP to HAp in organic media was examined in detail. In this study, hydrolyses of ot-TCP in a series of aliphatic alcohols were carried out. EXPERIMENTA L PROCEDUR E ot -TCP was provided by Taihei Chemical Industrial Co. Ltd. Hydrolyses of a -TCP in a series of aliphatic alcohols ( ethanol, 1-butanol, 1-hexanol, and 1-octanol ) were carried out ; the mixture of ot-TCP (0.01 mol) and 0.1 M ammonium aqueous solution (36ml) in aliphatic alcohols ( 50ml) was stirring for 2^^120 hours at 70 C. The initial pH value was adjusted to about 11.0 241

242


with ammonium solution. The reaction products were filtered off, washed with distilled water, and dried in air at 40 C for 5 hours. The obtained products were identified by X-ray diffractometry ( XRD ; Rigaku Geigerflex RAD I A ). The microstructures of HAp were observed by the scanning electron microscopy (SEM ; Hitachi FESEM H800) and transmission electron microscopy (TEM ; Hitachi H8000, 200kV). RESULT S AND DISCUSSIO N Hydrolyses of a -TCP in each aliphatic alcohols were carried out. Except that of a -TCP in ethanol, the reactions proceeded biphasically. Fig. 1 shows the X-ray diffraction patterns of the products ( reaction time ; 4 hours, reaction temperature ; 70 C ). In this reaction condition otTCP was partially transformed to HAp. Thus, the peaks corresponding to a -TCP and HAp were observed. Based on the relative intensity of the peak for a-TCP and HAp, the rates of transformation were compared. With increasing the hydrophobicity, the rates of hydrolyses

(0

c o

IE

CO

25

30

35

26 / degre e Figure 1. The X-ray diffraction patterns of the products prepared in each aliphatic alcohols (reaction time ; 4 hours, reaction temperature ; 70 C). a ; absence of alcohol, b ; 1octanol, c ; 1-hexanol, d ; 1-butanol, e ; ethanol, ; HAp, O ; a-TCP.

Transformationof a-TCP to Hydroxyapatite in Organic Media: K. Sakamoto et al.

30 /i m

243

1 lim

Figure 2. The SEM (left) and TEM ( right) photographs of a -TCP ( a ), HAp prepared in the absence of alcohol ( b ), and HAp prepared in ethanol ( c ) and 1-octanol ( d ).

244


increased as the following sequence ; 1-octanol > 1-hexanol > 1-butanol > ethanol. Even if at 40 C, Q:-TCP were partially transformed to HAp in these solvent systems. In 1-octanol and 1-hexanol, the complete conversions of a-TCP to HAp were achieved for stirring 24 hours at 70 C. The transformation rate in 1-octanol was compatible to that in the alcohol-free system. In 1-butanol ot -TCP was completely transformed to HAp after stirring for 72 hours at 70 C. However, in hydrophilic alcohol (ethanol) the complete conversion was not achieved after 96 hours at 70 C. Fig. 2 shows the SEM and TEM photographs of ot-TCP, and HAp prepared in each aliphatic alcohols. The particles of a-TCP were the smooth surfaces and irregular form. On the other hand, the figures of obtained HAp were a fiber like. The microstructures of HAp were influenced by hydrophobic or hydrophilic of aliphatic alcohols. HAp prepared in ethanol and 1-butanol were as well as that in the absence of alcohols. They were the mixture of platelet-form ( width ; 0.5 U m ) and fine needles ( length ; about 1 M m ). The products in 1-octanol and 1-hexanol were the needle-like particle (length ; 1.0^2.0 U m, width ; 0.1 M m ). Solubilities of the aliphatic alcohols in water at 70 C are as follows : ethanol ; -^, /1 / / /

-

~j

V"

’

PH 3

49 4 2

t :yo.^

Ol^ L

,

,

,

50

47 ,

Figure 3. Evolution of the ^’Ca quantity on the solid sample and in the host solution (pH 3, no stirring, room temperature)

pH6 PH 7 pH8 1

, 200

,

H

300

1

i

350

400

Ttane (hours)

Figure 4. Calcium dissolution in Ringer’s solution under different pH (no stirring, room temperature)

RESULT S AND DISCUSSIO N 1. Dissolution behaviour Firstly, we have tested the feasibility of the method. No chemical and morphological alteration caused by the irradiation has been noted by SEM and X-rays. After irradiation, dissolution tests were performed: the quantity of Ca on the solid sample decreases whereas the radioactivity in the liquid increases proportionally (figure 3). As an example, the dissolution under different pH (2 to 8) at room temperature was measured (figure 4). The group of curves show the strong influence of the pH on the kinetics. For the pH 2 and 2.5, after 2 or 3 days, the metal is completely bare with a total disappearance of coating. On the contrary, the more the pH increases, the more the dissolution slows down. At pH 8, the threshold of the spectrometer is not reached: the "^^Ca quantity is too low in the liquid. On each curve, a plateau appears: the dissolution phenomenon decreases. But a saturation doesn’t happen because any precipitation is observed on our samples. It is well known that plasma-spraying process strongly affects HAP coatings. Even though the CaP powders are rather stable, the coatings aren’t [4]. The plasma-spraying modifies the cristallinity, the specific surface and mainly the composition phases: with HAP, new CaP appear and some deshydroxylations happen. Moreover, the coating structure is favorable to a dissolution because of its lamellar aspect with many cracks, pores where some similar conditions to crevice corrosion are created (low pH, few or no fluid movement): the acid attacks are favoured. It is proved too, that some ions (for example Na^, CO3 ", in the Ringer’s solution) increase the solubility of HAP coatings [5]. The pH influence and some mechanisms (like the PO4 " ion complexations) [6] begin to be proposed for all these observations. Thermal treatment 500 C

|

Distance without 900 C (mm) 85 11.2 9.5 26.5 100 10.4 18.7 34.2 120 8.9 17.7 30.1 Table 2. Adherence (MPa) of HAP coating with the projection distance (mm) and post plasma-spraying treatment (3h30)

Eoating composition ’ HAP alone 25%Al203/75%HAP 50%Al2O3/50%HAP 75%Al203/25%HAP Alumina alone

Adherence (MPa) 11 11.6 12 12.6 15

Table 3. Adherence (MPa) according to the AI2O3 content in the coatings

278

Bioceramics Volume10 Ln<M> - (tt in HB>) 4,1

^

3,0 3,7 3,5

p^’

3,3 3,1 2.9 2,7 2.5

(

4,5

1

4.

5

1 S,5

1 6

1

0«^«Al2O3 25^*AI20J > 50 tetracalcium phosphate (TTCP) > mixture of TTCP and DCPA, i.e.,powder phase of CPC > crushed set c-CPC. It should be noted that crushed c-CPC also showed an inflammatory response even though it is an apatitic mineral. We concluded, therefore, that CPC shows excellent tissue response only when it is set to form apatitic mass. Thus, CPC should be used so that its setting reaction can be assured. INTRODUCTIO N Calcium phosphate cement invented by Drs. Brown and Chow consists of an equimolar mixture of tetracalcium phosphate (TTCP: Ca4(P04)20) and dicalcium phos› phate dihydrate (DCPA: CaHP04) [1-3]. When mixed with an aque› ous solution, it sets to form hydroxy apatite (HAP: Cal0(PO4)6(OH)2), the putative mineral of tooth and bone. The set mass shows excellent tissue re› sponse towards hard and soft tis› sues. However, conventional CPC (c-CPC) caused a severe inflamma› tory reaction when the paste, not the set mass, was implanted

Figure 1. 301

A4)pearance of rat abdomen 1 week after implantation.

302


subcutaneously in a rat immediately after mixing (Figure 1). In contrast fast-setting calcium phosphate cement (FSCPC) [4,5], which set within approximately 5 min, as opposed to 30 60 min - the setting time of c-CPC, and its anti-washout type (aw-FSCPC) [6,7], the paste would not be washed-out but set within approximately S min even if the paste was immersed in serum immediately after mixing, showed excellent tissue response. c-CPC was found to be completely crumbled whereas FSCPC and aw-FSCPC kept the same shape as at implantation. In addition, unreacted DCPA was found only in the case of c-CPC even 24 hours after implantation. One of the reasons for the inflammatory response observed in c-CPC may be the crumbling property of c-CPC. When the paste was arushed into powder, TTCP and DCPA could hydrolyze to form HAP, not in a symmetric way. It also forms HAP powder instead of the set HAP mass. In this investigation, several calcium phosphate powders, 1) TTCP, 2) DCPA, 3) mixture of TTCP and DCPA, and 4) crushed set CPC, were unplanted subcutaneously in rats and tissue response to each powder was examined to shed some light on the cause of the inflammatory reaction of c-CPC. MATERIAL S AN D METHOD S Specimens preparalion TTCP was made by heating the mixture of DCPA and CaCQ3 at 1500 C for 12 hours and crushed into powder as descr9)ed previously. DCPA obtained commercially was grounded in 90% ethanol to reduce the size to 0.9 (Jim in diameter. The powder phase of CPC was made by mixing an equimolar amount of TTCP and DCPA. The CPC powder was mixed with distilled water, at a powder to liquid (P/L) mixing ratio of 3.5 and kept in an incubator at 37 *C and 100 % relative humidity for 24 hours to get set mass. Some of the set mass was crushed to get a crushed set CPC. Animals and implantation procedure Ten-week-old male rats of Wistar strain obtained conmiercially (Charles River, Yokohama, Japan) and given standard pellets and water ad libitumywere used for the unplantation study. All powders were tested by implantation in all of the 20 rats. The rats were anaesthetized by Lp. injection of sodium pentobarbital (Nembutalfi, Abbott Co., Chicago, IL). For the implantation of CPC, the abdomen of the rat was shaved, washed and disinfected with iodine. Three longitudinal incisions of about 1 cm were made through the full thickness of the skin. Subsequently, lateral incisions to the subcutaneous pockets were created by blunt dissection with scissors. Each experimental material (0.3g) was unplanted using a cylindrical mold made by cutting thefrontportion of a 1 cm^ plastic syringe (Terumo, Tokyo, Japan). Set CPC was also implanted subcutaneously in rat as control materials. Finally, the wounds were carefully closed. Histological preparations At the end of the implantation period the rats were killed with an excess dose of Nembutal. After soft x-ray photographs were obtained to record the behaviour of calcium phosphate powers in each rat, the implant materials, including all surrounding tissues, were removed, fixed in 10% neutral buffered formalin and onbedded in methylmethacrylate (HistoDuifi, Leica Co., Nussloch, Germany). After polymerization, thin serial sections were cut using a rotary microtome. The sections were stained with hematoxylin-eosin and investigated by light microscopy. RESULT S

Mechanism of InflammatoryReaction of ConventionalCalcium Phosphate Cement:K. Ishikawa et al.

Figure 2 shows the typical appear› ance of the rat*s abdomen 1 week af› ter surgery. As shown, the most severe swelling with fluctuatio n by palpation was apparent around the DCPA . TTCP , the other compo› nent of the CPC , also showed an inflammatory response . The size of swelling was smaller in the case of the mixture of TTC P and DCPA , i.e., the powder phase of CPC . Crushed set CPC also induced an inflammatory response but its swelling was smallest in size. No gross evidenc e of inflammatory^ re› sponse was observe d when set CPC was implanted subcutaneousl y in rat (data not shown). The size of swelling formed 1 week after im› plantation was in the order of the following (Table 1)

303

^h^^’t^ Crushe d se t CPC Mixture of TTCP and DCPA

..-^ , -Figure 2. Appearance of rat abdomen 1 week after implantation.

DCPA > TTCP > mixture of TTCP and DCPA > crushe d se t c-CP C When the skin covering the calcium phosphate powder was cut with scissors, a copious inflammatory effusion compose d of serous, slightly viscous, yellowish, transparent fluid, was observed. In contrast, no effusion was observe d where set CPC had been implanted, and the set mass was covered only by a thinfibrouscapsule. Histological evaluatio n reveale d that a large vesicle containing abundant inflammatory effusion was formed around the calcium phosphate powder one week after implantation . The wall of the vesicle was compose d of thick vascular granulation tissue which containe d many foreign-bod y giant cells and moderate infiltration of inflammatory cells - consisting of lymphocyte s and plasma cells. In the cytoplasm of the foreign-body giant cells, calcium phosphate powder stained with hematoxyli n was frequently observed. Table 1 Implanted calcium phosphate and size of swelling, the amount of inflanunatory effusion formed subcutaneousl y in rat 1 week after implantation Particle

Powder Size^) dim)

DCP A TTC P Mixture of TTC P and DCP A Qushed set CPC

0.9 11.0

-

n.d.

Size of the Swelling^) Inflammatory Effusion) Length x Width x Height

28 21 18 14

X X X X

a) Average diamete r of calcium phosphate b) In mm effusion

16 X 8 x 11 X 9 x

8 4 2 1

Very Large Large Medium Small

c) Relative amount of inflammatory

304


DISCUSSIO N This investigation clearly demonstrated that TTCP and DCPA caused an inflammatory response even though they are the components of CPC. Also, crushed set CPC caused an inflanmiatory response even though it is the apatitic mineral. These resuhs are consistent with the results that c-CPC caused an mflammatory response when implanted immediately after mixing. It is confirmed that the important factor in obtaining satisfactory clinical results is to use CPC so that it can not be crumbled in the implanted area, CPC set to form HAP as shown in eq 1. However the reaction ]»roceeds in a symmetric way only when both TTCP and DCPA stay close. When TTCP aoid DCPA are apart for some reason, for example crumbling property, TTCP and DCPA hydrolyze to produce Ca(0H)2 and H3P04, respectively. The pH of the surrounding area will increase and decrease due to Ca(0H)2 and H3PO4 formation, respectively. 2Ca4(P04)20 + 2CaHP04 3Ca4(P04)20 + 3H20 10CaHPO4 + 2H2O

• •

•

Caio(P04)6(OH)2

Caio(P04)6(OH)2 + 2Ca(OH)2 Caio(P04)6(OH)2 + 4H3P04

(1) (2) (3)

It should be noted that crushed set CPC also caused an inflammatory response even though the size of the swelling was smallest within the experimental group. These results indicate that the size of the implant materials is an important factor to decide the tissue response. The largest size of swelling and the largest amount of inflanunatory effusion observed in DCPA may be the resuU of two factors. First, hydrolysis of DCPA produces H3P04 as by product and reduces the pH of the surrounding tissue. Second, the particle size is smallest among the calcium phosphate powders examined in this present study. These factors are thought to be oweed, in part, to the inflammatory response of c-CPC since unreacted DCPA was found in the case of cCPC. Further evaluation of tissue response to calcium phosphate powder with controlled particle size is awaited to understand the factors to determine tissue response. SUMMAR Y All calcium phosphate mineral relating to the component of CPC mcluding apatite showed inflammatory response when they are unplanted in powder form. Therefore, CPC should be used so that its setting reaction can be assured. ACKNOWLEDGMEN T This investigation was supported in part by a Grant-in-Aid for Scientiflc Research from the Ministry of Education, Science, Sports and Culture, Japan, and in part by a Grant-m-Aid for Scientiflc ResearchfromUehara Memorial Foundation.

REFERENCES 1. 2. 3. 4. 5. 6. 7.

Brown, W.E. and Chow, L.C. US Patent No. 4,612,053 1986. Brown, W.E. and Chow, L.C. In: Cements Research Progress, American Ceramic Society, Westerville 1986, 351-379. Chow, L.C. and Takagi, S. In: Specialty Cements with Advanced Properties, Materials Research Society, Pittsburgh 1989, 3-24. Ishikawa, K., Takagi, S., Chow, L.C. and Ishikawa Y. J Mater Sci: Mater Med 1995, 6, 528-533. Miyamoto, Y., Ishikawa, K., Fukao, H., Sawada, M., Nagayama, M., Kon, M. and Asaoka, K. Biomaterials1995,16, 855-860. Ishikawa, K., Miyamoto, Y., Kon, M., Nagayama, M. and Asaoka K. Biomaterials1995, 16, 527-532. Miyamoto, Y., Ishikawa, K., Takechi, M., Yuasa, M., Kon, M., Nagayama, M. and Asaoka, K. Biomaterials1996,17, 1429-1435.


FRACTUR E OF ALUMIN A CERAMI C HEA D IN TOTA L HI P ARTHROPLAST Y -REPOR T OF TW O CASE S WIT H HISTOLOGICA L EXAMINATIO N AND PARTICL E CHARACTERISATION Y. Kadoyal, A. Kobayashi^ P. A. Revell 2, H. Ghashi 1 ,Y. Yamano^, G. Scott ^ and M. A. R. Freeman^. 1 Dept. of Orthopaedic Surgery, Osaka City University Medical School. Osaka. Japan. 2 Osteoarticular Research Group, Dept. of Histopathology. Royal Free Hospital School of Medicine. Pond Street, London NW3 2QG. UK. 3 Bone & Joint Res. Unit. Royal London Hospital. Ashfield Street, London El IAD. UK.

ABSTRAC T Two fractures of the ceramic femoral head are reported. Detailed histological examination and SEM characterisation of the extracted particles were performed. Fragmented ceramic acted as a third body which caused severe metal and polyethylene wear, so that urgent revision procedure should be indicated. SEM and histological study showed that ceramic particles themselves were small enough to elicit foreign body reaction. It was also suggested that hydroxyapatite coating might prevent particle migration and subsequent osteolysis. KE Y WORDS : ceramic fracture, hydroxyapatite coating

total hip arthroplasty,

metallosis, ceramic particles,

INTRODUCTIO N The use of a ceramic femoral head in total hip arthroplasty (THA) has been popular because it produces much less polyethylene wear compared with a conventional metal head[1.2]. Ceramic has been preferentially utilised in younger and active patients where the reduction of wear is of particular importance to prevent osteolysis in the long term. However, this material is extremely hard and brittle thus susceptible to fracture. Although several fracture cases of ceramic head has been reported [3-9], the exact nature of the ceramic particles and consequent histological reaction to this particle has not been well documented. In this paper, two failures of the ceramic femoral head were investigated with detailed histological examination of the periprosthetic tissues. Furthermore, particles were extracted by a tissue digestion method [10,11], and characterised with scanning electron microscopy (SEM) . CAS E 1: A 35 year old man who had ankylosing spondylitis underwent left THA in 1988. The component was a Freeman hydroxyapatite (HA) coated stem (Corin Medical Ltd. Cirencester, UK) with 26 mm alumina ceramic head (Vitox_, Morgan Matroc Ltd. Surrey, UK). The acetabular component was HA coated superolateral fin (SLF) design. Thirty-one months after the operation, the patient developed sudden left hip pain when lifting a heavy load. A radiographic examination after 1 month demonstrated a fracture of femoral head (Fig. 1). At revision, the ceramic head was shattered into multiple small fragments. Marked metallosis was noted and the surface of the Morse taper was severely abraded and roughened (Fig. 2). The polyethylene liner had been deeply scored by the trunnion of the femoral component (Fig. 3). The interfaces of both components were sound and well fixed. Histologically, there was good bone growth onto the HA coating over much of the stem. In spite of a heavy metal and ceramic particle deposition in the bone marrow (Fig.4,5), there was no clear evidence of heavy infiltrate immediately next to the HA coating(Fig.4). 305

306


SEM examinatio n on the extracte d particles showed that the majority were submicron metal particles. However, ceramic particles around 5 ^m in diamete r were also observed (Fig.6). CAS E 2: A 57 year old woman who had osteoarthriti s was managed with right THA in 1991. The component configuratio n was the same as the first case. In 1995, the patient felt discomfort in her hip and radiography demonstrate dfi-actureof the femoral head. There was no evident history of trauma to the hip.

Figure 1. Fractured ceramic femoral head (white arrow).

Figure 2. Fragmented ceramic Figure 3. Deeply scored head and damaged taper. polyethylene cup (arrowheads).

Figure 4. Metal particles (*) in the bone marrow . HA-bone interface ( arrowheads ) remains intact. x50

Figure 6 SEM photograph showing small ceramic particles (arrows), bar =lpm

Figure 5. Ceramic particles (arrows) exist between the band of metal particles (*). x200

Figure 7. Ceramic particles in the tissue (arrows), Smaller particles are also present. x2(X)

Fracture of Alumina Ceramic Headin TotalHipArthroplasty: Y. Kadoya et al. 307

The intraoperative findings at revision surgery were identical to the first case including diffuse metallosis and damaged femoral taper and polyethylene. The acetabular component was loose. Histologically, numerous small particles presumed to be ceramic particles (2-5|Lim) were observed (Fig.7) with severe metallosis. Polyethylene particles were occasionally seen and these were generally very large (50-100)im) and predominately provoked a giant cell reaction. DISCUSSIO N The use of a ceramic femoral head has been advocated in the young and active patient because of its improved wear characteristics when articulated with polyethylene. However, due to its extreme hardness and brittleness, several cases of fracture have been reported[3-9]. One common findinjg among these fracture cases is the existence of severe and diffuse metallosis. Following ceramic head fracture, small ceramic fragments embedded in the plastic acetabular component are potentially one of the most abrasive materials[3]. The severe damage on the trunnion in our cases has confirmed that fragmented ceramic could act as a third body which causes severe metal wear necessitating urgent revision surgery . In the previous literature on fracture of the ceramic head, few papers mentioned the existence of ceramic particles probably because of severe metallosis. Consequently, no attempt was made to extract and characterise the particles in the tissue. We demonstrated histologically that there were abundant ceramic particles in the infiltrating granulomatous tissue. Furthermore, the extraction and SEM study confirmed that these particles were around 5 ^m which is in accord with the histological observation that they were small enough to elicit a foreign body reaction. Although the duration of particle challenge was relatively short, it was suggested that HA coating might prevent particle migration to the bone-implant interface and subsequent osteolysis. SUMMAR Y This paper highlighted the role played by the fragmented ceramic head as the third body which accelerate the wear of metal and polyethylene. The exact size of the ceramic particles was determined (^ Sîm) and shown to be small enough to elicit foreign body reaction. HA coating acted as a seal against particle migration at least during the observed period. REFFERENCES 1. Oonishi, H., Takayama, Y., Clarke I.C, and Jung, H. J. Long-Term Effect of Medical Implants. 1992 2, 37-47. 2. Davidson, J. CUn. Orthop. 1993, 294,3 61-378. 3. Kempf, I and Semlitsch M. Arch. Orthop. Trauma Surg. 1990, 109, 284-287. 4. Hummer, CD., Rothman, R.H., Hozack W.J. J. Arthroplasty. 1995, 10, 848-850. 5. Higuchi, F., Shiba, N., Inoue, A. and Wakebe, I. J. Arthroplasty. 1995, 10, 851-854. 6. Callaway, G.H., Flynn W., Ranawat, C.S. and Sculuco T.P. J. Arthroplasty. 1995, 10, 855859. 7. Krikler, S., Schatzker, J. J. Arthroplasty. 1995, 10, 860-862. 8. Michaud R.J., Rashad, S.Y. J. Arthroplasty. 1995, 10, 863-867. 9. PuUiam I.T. and Trousdale R. T. J. Bone and Joint Surg. 79-A, 1997, 118-121. lO.Campbell P.,. Ma, S., Yeom, B., McKellop, H., Schmalzried, T,P. and Amstutz, H.C. J Biomed. Mater. Res. 1995,29,127-31. ll.Kobayashi, A., Bonfield, W., Kadoya, Y., Yamac, T., Freeman, M.A.R., Scott, G., Revell, P. A. Proc Instn Mech Engrs, Part H in press.



EXPERIMENTA

L COMPARATIV E S T U D Y BETWEE N

ROUGH-BLASTE D AN D HYDROXYAPATIT Young Ho Kim\ Tae Soo Park^

Jong Seok Park^

II Yong Choi\

E COATE D IMPLANT S Myung Ryool Park\

Department of orthopedic surger>^ Hanyang University, Kuri Hospital, 249-1, Kyomoon-Dong, Kuri, Kyunggi-Do, 471-020, Korea Department of orthopedic surgery, Soonchunhyang University, Chonan Hospital, 23-20, Bongmyung-Dong, Chonan, Chungchungnam-Do, 330-100, Korea ABSTRAC T We performed radiographic, biomechanical and histologic comparative assessment between rough-blasted (KB) and HA coated implants to identify the efficacy of HA coating on roughblasted titanium compared to KB surface on titanium in dogs. The results were as foUowings. Radiographically, HA coated implants had earlier and more abundant incorporation and proliferation of bone. Biomechanically, push out failure load increased gradually until 1 year after implantation in both groups and was significantly higher in HA coated implants compare to RB implants since 3 months after implantation. Histologically, more than 90% of surface coverage by bone was achieved since 3 months after implantation in HA coated implants but such same level of surface coverage was achieved as late as 1 year after implantation in RB implants, and earlier and more profuse production of osteoid and earlier maturation of bone with less intervening fibrous tissue around implants were found in HA coated implants compared to RB implants. Conclusively, more active osteoconduction with profuse surrounded marrow tissue maintained until 1 year after implantation in HA coated implants compared to RB implants. INTRODUCTIO N In experimental study, HA coated implants show better bone ongrowth than plain titanium press-fit or porous coated implants[2, 6] and HA can help to achieve such ingrowth even under condition of micromotion[4, 5]. But there were some problems such as the possibility of osteolysin as a reaction to loose HA particls and delamination[l]. We performed radiographic, biomechanical and histologic comparative assessment between rough-blasted(RB) and HA coated implants to identify the efficacy of HA coating on roughblasted titanium compared to RB surface on titanium in dogs. MATERIAL S AN D METHOD S Cylindrical rod of Ti-6Al-7Nb titanium alloy were prepared, measuring 4.5mm in diameter and 6mm in length. Two types of coating were applied on the rod, which were HA coating with 5 //m of thickness using plasma spray technique and rough blasted coating. The rods were inserted into predrilled holes in the lateral cortex of 309

310


adult canine femur using sterile surgical techniques. The holes were drilled slightly oversized (4.7mm), which allow the rods to be implanted without undue laxity. A total 42 rods were inserted into the femur of 7 dogs. 21 RB rods were implanted on one side of femur and 21 HA coated rods were implanted on the other side of femur without surgical complication. Dogs were sacrificed 2 at 6 weeks, 2 at 3 months, 2 at 6 months and 1 at 1 year after implantation. 42 bone segments containing the plugs were obtained after cutting with air saw. Radiographs were taken for all 42 segments taken each time to identify the radiographic difference for the osseointegration of 2 kinds of implants with time. The 34 bone segments containing the plugs were positioned in a testing jig to allow accurate alignment of the loading axis to the long axis of the plugs. The plugs were pushed out from the surrounding bone using an Instron machine 8501 with a crosshead speed of 1 mm/minute to get push out failure load each time. The 8 bone segments which were selected at random each time were prepared with the section of 100 lim thick for light microscopic examination. Villanueva stain was used for each prepared section. Percent surface coverage was estimated by the use of a transparent square grid [7]. RESULTS Radiographically, HA coated implants had earlier and more abundant incorporation and proliferation of bone. Biomechanically, push out failure load increased gradually until 1 year after implantation in both groups and was significantly higher in HA coated implants compare to RB implants since 3 months after implantation(Table 1). Histologically, more than 90% of surface coverage by bone was achieved since 3 months after implantation in HA coated implants but such same level of surface coverage was achieved as late as 1 year after implantation in RB implants(Table 2), and earlier and more profuse production of osteoid and earlier maturation of bone with less intervening fibrous tissue around implants were found in HA coated implants compared to RB implants. But histologic findings at 1 year after implantation were similar in both groups, showing well surrounded mature bone, except more profuse appearance of marrow tissue around HA coated implants(Fig. 1, 2, 3, 4). Table 1. Pushout failure load Postop time 6Wks 3Mons 6Mons lYr

Failure load(N) No.

RB

5 423.60–84.83 5 613.13–151.28 4 652.61 –133.07 2 982.47 –115.84

P-value HA

489.32–83.73 791.67–126.89 832.18–59.84 1178.67 –125.56

0.055 0.012 0.018 0.022

ComparativeStudy BetweenRough-Blasted and Hydroxyapatite Coated Implants: Y.H. Kim et al.

311

Table 2. Percent surface coverage RB(?’6)

HA(%)

6 Wks 3 Mons 6 Mons 1 Yr

30 45 60 95

70 90 95 100

DISCUSSION AND CONCLUSION In our biomechanical study, push out failure load was significantly higher in HA coated implants comared with RB implants, which finding was similar to other reports[3, 7]. Our results represented exceptionally high osseointegrative properties in HA coated implants until at least 1 year after implantation, especially since 3 months after implantation. The findings that more than 90% of surface coverage by bone was achieved as early as since 3 months after implantation in HA coated implant may make a significant contribution to the mechanical strength of the interface between bone and HA. In our experiments, dogs were allowed to move freely immediately after implantation, which means they were under condition of micromotion. Even under such condition of micromotion, earlier and more profuse production of osteoid and earlier maturation of bone with less intervening fibrous tissue were achieved in HA coated implants compared with RB implants, which findings were supported by the other reports[l, 4, 5]. Even if two kinds of

Fig 1. Postoperative 6 weeks histologic section (Villanueva stain ; original magnification x 1) of rough-blasted plugs implant in bone shows scanty distributed osteoid around the implant.

Fig 2. Postoperative 6 weeks histologic section (Villanueva stain ; original magnification x 1) of hydroxyapatite coated plugs implant in bone shows abundant distributed osteoid around the implant.

312


Fig 3. Postoperative 1 year histologic section (Villanueva stain original magnification X 40) of RB plugs implant in bone shows well maturated bone in close contact with the implant associated with relatively poor surrounding marrow component.

Fig 4. Postoperative 1 year histologic section (Villnueva stain original magnification X 40) of HA coated plugs implant in bone shows well maturated bone in closer contact with the implant associated with abundant surrounding marrow component, compared with the findings of RB implant.

implants had similar histologic findings at 1 year after implantation, which were well surrounded mature bone around the implants, normal morrow tissue existed more abundantly around HA coated implants compared with RB implants. The findings may suggest osteoconductive activity of HA coating was still maintained and remained until 1 year after implantation and efficacy of HA coating for the capacity of osseointegration after implantation. Conclusively, More active osteoconduction with profuse surrounded marrow tissue maintained until 1 year after implantation in HA coated implants compared to RB implants. REFERENCE 1. Bleobaum R.D. and Dupont J.A., / Arthroplasty, 1993, 8, 195-202. 2. Geesink R.G,T. and Hoefnagels N.H.M., / Bone Joint Surg, 1995, 7 7 - B , 534-547. 3. S(5balle K., Acta Orthop Scand suppl 1993, 64, 1-58. 4. S0hdl\e K., Brockstedt-Rasmussen H. and Hansen E.S., Acta Orthop Scand, 1992, 63, 128-140. 5. S(Pballe K., Brockstedt-Rasmussen H., Hansen E,S., / Bone Joint Surg, 1993, 75-B, 270-278. 6. SOdlXe K, Hansen E.S., Brockstedt-Rasmussen H. and Pedersen CM., Acta Orthop Scand, 1990, 61, 299-306. 7. Wong M., Eulenberger J Schenk R. and Hunziker E., / Biomed Mat Res, 1995, 29, 1567-1575.


MECHANICA L AN D BIOLOGICA L PROPERTIE S OF ALUMIN A BEA D COMPOSIT E M. Kobayashi^ ^ T. Nakamura^ T. Kikutam^ Y.Okada^ N. Ikeda^ S. Shinzato^ aiKiT.Kokubo^ ^Dq)artment of Orthopaedic Surgery, Faculty of Medicine, Kyoto University, Kawahaia-cho 54, Shogoin, Sakyo-ku, Kyoto 606-01, Jân, Ôrttiopaedic Surgery, Otsu Red Cross Hospital, Otsu, Japan, ^Nippon Electric Glass Co. Ltd., Otsu, Jân, and ’*Dq)artment of Material Chemistry, Faculty of Engineering, Kyoto University, Kyoto, Japan. ABSTRAC T We have developed a new composite (ABC) consisting of alumina bead powder and bisphenol-^glyddyl methacrylate(Bis-GMA)4)asedresin, which has both high mechanical strength andexcellentosteoconductivity. Aluminabeadof99.7%puretooksphericalforms // minaverage size andcontainedamorphous,and 5-and r-crystalphases. Anothercon5)osite(SGC)filledwith amorphous silica was used as referential material. The proportion offilleradded to each composite was 70% w/w. Mechanical testing of ABC indicated that it would be strong enough for use under weight-bearing conditions. Histological examination using rat tibiae forup to 26 weeksrevealedthat ABC had excellent osteocondactivity, which was eqaivalent to that of a con5)osite containing AW-GC rq)orted previously. And at 26 weeks, no marked biodegradation had occurred. Whereas,in SGC-implanted tibiae, there was poor cJrect bone formation even at 26 weeks. ABC may have a potent promise as a both mechanically strong and highly biocon5)atible material. KEYWORDS ; Alumina, Bis-GMA, Composite, Osteoconduction, Biological Property INTRODUCTIO N Alumina ceramics have good biocompatibility, high mechanical strength, high resistance to fatigue, and excellent lubrication properties. [1] However, alumina ceramics arebioinert andhaveno bone-bonding ability. Thus, attenq)ts havebeen madeto in5)rove the bone-bonding strength, mainly by surface modification. We have succeeded in developing a new conq)osite (ABC), consisting of Bis-GMA-basedresin as an organic matrix and alumina bead powder, produced by fusing a -alumina powder and subsequently (|iencfaing it as an inorganic filler, and in demonstrating that ABC has excellent osteooonductivity.[2] hi order to develop a material which has high bioactivity, higji mechanical strength, and less potential for fatigue, osteooonductivity in rat tibiae of ABC for iq) to 26 weeks after in^lantation were assessed. MATERIAL S AN D METHOD S 1) Prq)aration of powder Bulk a -alumina (99.5%pure AI2O3) was prq)aredin an electrical melting fmnacefrom calcined alumina powder,producedbyBayer’sprocess. The bulk a -aluminawasthenpulverizedandparticles under 10 fi m in diameter were coUected. The coUectedpowderwas fused and quenched subsequeiitily to produce alumina bead powder (AL-P). Powder XRD andFT-IRRS oftheAL-P showed that it 313

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contained amorphous and 6 -crystal phases of alumina in its main crystal structure . Its average paitide size was 3.0 fi m, the density was 3.6 g/cni, and specific surface area was 0.7 t i i / g. Spherical particles of amorphous silica glass powder (SG-P) were also prq)aredby the fusingcpendiing method firomhighly purified quartz (>99.7% pure). It had the density of 2.2 g/cni, average particle size of 3.0 /x m, and specific surface area of 12 mV g. Each powder was silane4reate d with r -methacryloxypropyltrimethoxysilane , and benzoyl peroxidB at 0.4% per unit weight of the treate d powder was added Bis-GM A and triethyleneglyoo l dimethacrylat e (TEGDMA ) weremixed in equal weight. N,N-dimethyl^-toluidine , at 1.0%perunit weight ofthemixtur e ofBis-GM A and TEGDMA , was dissolved.[3-6 ] The weight ratio of the filler powder mixed into the composite was 70%. The composite was prepared by mixing the ^propriate powder into the mixture of Bis-GM A and TEGDM A and stirring it for 1 minute. It was polymerize d within 3-4 minutes. The two types of composite , containing either AL- P or SG-P as afiller,were designate d AB C and SGC , respectively . The ultimate con5)ressiv e strength, bending strength, dastic modulus of bendng (Young’s modulus), tensile strength and fiacturetoughness of ABC and SGC , measured after soaking in simulatedbocf y fluid (SBF) at 37t: for 1 day, were 196–4 and237– 14 (MPa), 151 – 10 and 157 –10(MPa),7.2–0.2and8.6–0.2(GPa),58–3and59–5(MPa),andl.44–0.05andl.69–0.1 1 (MP a m^ ^), respectively . [2] 2) Aiumal e>q)erimen t Ten-week-ol d male Wistar rats were operate d on under general anesthesi a (Nembutal: 40 mg/kg hodyweight). Cortical bone defect s measuring 2 X 5 nun were aeated on the medial aspect of the proximal metaphysis of both tibiae, and the bone marrow was curette d The intramedullary canals firom both bone defect s in each individual animal were packed with the same kind of composite , and allowed to cure in situ.Twdve rats receive d ABC and twelve SGC ^ with three rats in each groi?) being killed at 2, 4, 8, and 26 weeks after the operation.[2-6 ] 3) Histological examinatio n All tibial segments containing composite samples were excise d and dehydrate d in serial dilutions ofethanol , then embeddedi n polyesterresin . Thin sections (500 // m thick) were cut with a band saw (EXAC T BS-3000, Nonderstedt , Gemiany), papendicula r to the axis of the tibia Two sectionsfiromeach tibia in the ABC and SGC (i.e. 12 specimen s in total for each subgroup) were ground to a thickness of 100 /x m using a diamond 1^ disk (Maruto Ltd., Tokyo, Jân)for contact microradiogr^hy. Several sectionsfiromthe six subgroups were ground to a thickness of 100 ix m for Giemsa surface staining. Several 500 /x m sections takenfiromthe AB C and S GC gjcovps at every time interval studied were polished with damond pêr. These sections were used to study the bone-compositeinterface,usin g a SEM (Hitachi S-800, Tokyo, Jân) connectedt o anEDX(Horib a EMAX-3000 , Tokyo, Japan).[2-6 ] RESULT S Histological examinatio n by contact microradiogr^hy and Giemsa surface staining reveale d that new bone had formed direcdy on the ABC surface, without an intervenin g soft tissue layer, by 2 weeks post-inq)lantation , and at 4 and 8 weeks, newly formed bone almost completel y surrounded the con9)osit e surface within the tibiaeimplanted with ABC . Furthermore,this was maintained at 26 weeks after the operation . However, in SGC-implanted tibiae,poor direct bone formation was observed on the SG C surface throughout the e5q)erimenta l period (Figure 1). Examination by SEM-ED X clearly demonstrate d direct bone formation on the AB C surface (Fig. 2a). ED X profiles of the bone-ABC interface reveale d slightiy inaeased intensity for calcium and

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.* .

^^) (b) Figure 1. Contact inicroradiogram of ABC and SGC in rat tibiaeat 2 weeks after implantation (a) AB C and (b) SGC . C; composite , B; bone.

(a) (b) , (a) Back-scattere d electro n image, (b) Figure 2. SEM-ED X of AB C at 8 weeks after implantation EDXprofiles. C, composite ; B, bone; Ca, caldum; P, phosphorus; Al, alununum.

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phosphorus (Fig. 2b). In SGC implanted tibiae, no such a layer was evident until 26 weeks. DISCUSSIO N The results of the present stu^ indicate that ABC has excellent osteooonductive ability. Although few studies have been conducted on amorphous alumina as a biometerial, amorphous alumina is thought to have excellent biocompatibility.[7] In the previous study, although apatite was not induced on the ABC surface after soaking in SBF for 28 days, the ABC surface had made direct contact with bone via a lêr containing calcium, phosphorus, and alumina powder. [2] Li et al. rq)orted that alumina gd did not induce âtite formation when inunersed in SBF for 21 d^s, whereas both pure silica gel and gd-derived titania were hydroxyâtite inducers. [8] However, once a material is implanted in the body, it elicits several responsesfromliving tissues. They include protein adsorption and cell attachment and adhesionas well as ionic exchange. We deduced that ABC has the ability to bond directly withbone, which was induced not by a simple diemical reaction but by some surface property of the AL-P whidi encouraged calcification or apatite formation due to the actions of proteins and cdls in vivo. The predse medianism of direct bone formation on the ABC surface is as yet undear. However, we consider ABCto showpromiseasabasisfordevelopingahighly osteoconductive and mechanically strong biomaterial. We are now plarming to evaluate its bone4)ondng ability, and mechanical properties after long-term in5)lantation. REFERENCES 1. Z. Li, T. Kitsugi, T. Yamamuro,Y. S. Chang, Y. Senaha,H. Takagi,T. Nakamura,arKi M. Oka, "Bone-bonding behavior under load-bearing conditions of an alumina ceramic implant incorporating beads coated with glass-ceramic containing apatite and woUastonite," J. BiomedMater.Res., 29, 1081-1088 (1995). 2. M. Kobayashi, T. Kikutani, T. Kokubo, and T. Nakamura, "Direct bone formation on alumina bead con5)osite," J. BiomedMater.Res., in press. 3. K. Kawanabe, J. Tamura,T. Yamamuro, T. Nakamura,T. Kokubo, andS. Yoshihara, "A new bioactivebone cement consisting of BIS-GMA resin andbioactive glass powder," J. Appl.Biomater.,4, 135-141 (1993). 4. J. Tamura,K. Kawanabe,M. Kobayashi,T.Nakamura,T. Kokubo, S. Yoshihara,andT. Shibuya. "Mechanical and biological properties of two types of bioactivebone cements containing MgO-CaO-Si02-P205-CaF2 glass and glass-ceramic powder," J. Biomed. Mater.Res., 3 0, 85-94 (1996) 5. M. Kobayashi, T. Nakamura, J. Tamura, H. lida, H. Fujita, T. Kokubo, andT. Kikutani, ’^Mechanical and biological properties of bioactive bone cement containing silica glass powder," J. Biomed.Mater.Res., in press. 6. M. Kobayashi, T. Nakamura, J. Tamura,T. Kokubo, andT. Kikutani, "Bioactivebone cement: con5)arison of AW-GC filler with hydroxyâtite and 13 -TCPfillerson mechanical and biological properties," J. Biomed.Mater.Res., in press. 7. A. Naji and M.F. Harmand, "Cytocompatibility of two coating materials, amorphous alumina and silicone carbide, using human differentiated cell cultures," Biomaterials,12, 690-694 (1991). 8. P. Li, C. Ohtsuki, T. Kokubo, K. Nakanishi,N. Soga, and K. De Groot, "The role of hydrated silica, titania, and aliunina in inducing apatite on implants," J. Biomed.Mater. Res., 2S, 1-15(1994).


REMODELIN G STEM S

OF BON E AROUN D HYDROXYLAPATITE-COATE

D

FEMORA L

A. A. Edidin and M. T. Manley R & D Corporate, Osteonics Corp., Allendale, NJ, USA

ABSTRAC T Stress shielding of the proximal femur following hip arthroplasty has been well documented around both cemented and cementless femoral stems. Attempts to limit the degree of stress shielding fall into two primary classes: Reduction of the structural stiffness of the stem, usually by reduction of its modulus of elasticity, and the addition of proximal interface enhancements designed to transmit axial forces as proximally as possible to the surrounding bone. We report on a series of patients who underwent THA using femoral stems at opposite ends of the design space. Specifically, one part of the cohort received an extensively porous-coated CoCr prosthesis while the other part received a proximal HA-coated Ti6A14V prosthesis. Using Dual Energy X-Ray Absorptiometry (DEXA) Bone Mineral Density (BMD) quantification techniques, we were able to detect an up to 30% greater retention of proximal bone density in the Ti6A14V cohort as compared to the CoCr cohort. INTRODUCTIO N Stresses in the femur arise from the axial and bending components of the load across the femoral head. While the ratio of axial to bending load is patient and gait pattern specific, the bending component dominates by about 3:1. Thus attempts to transfer more of the bending load to the bone must be focused on reducing the structural stiffness (EI) of the implant. While the moment of inertia is more quickly reduced than the modulus by simply narrowing the implant’s cross-section, physiology using cemendess implants dictates that an implant must be large enough to contact the endosteum if biological integration following mechanical stability is to occur. Thus reduction of the modulus.of elasticity (E) is the most effective means of reducing the implant’s structural stiffness. Attempts to transfer the axial load component proximally are generally limited to encouraging biological integration in the metaphyseal region using coatings or ingrowth regions. This report compares two successful femoral arthroplasty stems at opposite ends of the mechanical spectrum. The first stem is made of cobalt-chromium alloy with a modulus of 220 MPa, and is extensively coated using porous sintered beads. The second stem is made of Ti6A14V alloy with a modulus of 110 MPa and is proximally HA-coated. While both stems have a long (>10 yr.) and successful clinical record, the latter stem would be expected to stress shield the femur to a lesser degree. We used the DEXA technique to determine if in fact this expectation was met. 317

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MATERIAL S AND METHOD S Patients requiring unilateral primary arthroplasty of the hip with simple osteoarthritis as a diagnosis were eligible for inclusion. Inflammatory arthritis, trauma, femoral dysplasia, and trochanteric osteotomy were grounds for exclusion. Twelve patients received an extensively porous-coated CoCr hip stem (AML, Depuy, Warsaw, IN) and 18 patients received a proximally HA-coated Ti6A14V hip stem (Omnifit-HA, Allendale, NJ). The mean age of the patients in the first group was 54 years; patients in the second group had a mean age of 50 years. There were 5 males and 7 females in the CoCr cohort with 8 males and 10 females in the Ti6A14V cohort. DEXA films were obtained preoperatively, at five days post-operatively, and at six weeks, six months, and one and two years post-operatively. BMD was measured both in the Gruen Zones about the femur and alternatively in 2 cm intervals about the femur. The former measurements provided a stem-proportionalized breakdown of BMD changes, while the latter provided a stemlength independent assessment of BMD changes. RESULT S Radiographs of patients with each of the implants are shown in Figures la and lb. Both patients show good bone quality in keeping with their relatively young age and simple OA diagnosis.

Figure la: Lateral radiographof a patientwith the CoCr stem.

Figure lb: Lateral radiographof a patient with theTi6Al4V stem.

Bone mineral density changes for the anterior-posterior view Gruen zones are shown in Figure 2a and 2b for the HA-coated Ti6A14V and porous-coated CoCr stems respectively. In the proximal zones there was 25% greater bone resorption measured in patients receiving the CoCr stem as opposed to the Ti6A14V stem at two years follow-up. The bone resorption was also seen to move further down the stem into zones 2 and 6 in patients implanted with the CoCr stem.

Remodelingof Bone Around Hydroxyapatite-CoatedFemoral Stems: M.T. Manley et al.

Zone 1

Zone 7

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Zone 4

Post op S ix Weeks 0 6 Months H i Year [ 2 Year s

Figure 2a: AveragedBMD Changes by Gruen Zone as measuredin the patientcohort receivingtheHA-coated prosthesis.

Zone 1 0 -5 10 15 20 -25 30 -35 -40

Zone 7

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^ I I I’ I I I’ 111^ IM ’ HP’ LP I T 11 i^ i r I

91B

rj |^[r n

I

Post op S ix Weeks De Months H i Year 2 Year s

Figure 2b: AveragedBMD Changes by Gruen Zone as measuredin thepatient cohort receivingtheporous-coatedprosthesis.

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DISCUSSIO N Maintenance of proximal bone density at preoperative levels is not expected nor possible using existing technology. In order to preserve as much bone stock as possible, the procedure should still be designed to minimize proximal stress shielding. This study showed that by using a stem with a lower structural stiffness in conjunction with a biocompatible proximal coating, the degree of stress shielding could be substantially reduced. Because the femoral stems investigated in this study bound the extremes of the design space available today, the results presented herein may be considered to bracket the extremes of expected bone loss after hip arthroplasty. Thus the use of a lower-stiffness HA-coated stem may reduce the magnitude of proximal stress shielding by up to 25% at two years. In addition the region of most pronounced resorption is limited to the proximal two zones, as opposed to the mid-stem bone resorption seen using the CoCr porous-coated stem. ACKNOWLEDGMENT S The authors gratefully acknowledge the participation of William Jaffe, M.D., Fredrick Jaffe, M.D., and David Scott, M.D.


PROCESSIN G AN D CHARACTERISATIO N OF BIOLOGICA L HYDROXYAPATIT E DERIVE D FRO M CATTLE , SHEE P AN D DEE R BON E Michael R. Mucalol, Glenn S. Johnson 1 and Michel A. Lorier^ iChemistry Department, University of Waikato, Private Bag 3105, Hamilton, New Zealand, ^Meat Industry Research Institute of New Zealand, P.O. Box 617, Hamilton, New Zealand

ABSTRAC T In New Zealand, the routine slaughter of cattle, sheep and deer produces a large amount of waste bone which is normally converted to low-cost blood-and-bone fertiliser. The porous architecture of bone makes it a valuable material for use in biomedical implants and thus this poster will describe a study into the processing and characterisation of waste bone from animal species for conversion to materials for clinical purposes. Bone samples from these species were defatted using novel methods such as microwaving which was found to be a highly rapid fat removal method. Defatted bone cubes were then bleached using hypochlorite reagents. Infrared spectroscopy monitored bulk fat removal and have demonstrated that the hypochlorite treatment decollagenated bone cubes while leaving carbonate intact. Solid state NMR showed there was still some tenaciously held organic matter in the samples even after hypochlorite treatment. The work demonstrates that waste animal bone can be efficiently processed to produce modifiable materials for clinical use. KEYWORD S bone, hydroxyapatite, implants, defatting, deproteination, FTIR/NMR INTRODUCTIO N New Zealand as a major producer of meat foodstuffs often has to deal with the large amounts of by-products that result from the routine slaughter of livestock. Often these by-products have no use other than as materials for low value products or else are disposed of, which can potentially lead to environmental problems. A collaborative project between the University of Waikato and the Meat Industry Research Institute of New Zealand was established with the aim of converting one of these waste products, animal bone, to high value-added products. The unique porous architecture of bone makes it an ideal material for use in non-loading bioactive implants where tissue ingrowth is an important requirement. In addition, such bone is an obvious source of hydroxyapatite which may be used in the production of synthetic biomaterials or as a matrix for Drug Delivery Systems. An additional advantage is that the restrictions which necessitate the extraction of bone materials only from controlled herds (e.g. Kiel bone in Germany) in the European Community do not apply in New Zealand. This project was initially inspired by an actual clinical application which required that the bone material be aesthetically presentable and cuttable to a desired form. As-received bone after defatting is extremely hard and therefore cannot be cut or preformed into a convenient size as can synthetically produced porous hydroxyapatite. Further treatment of the bone was, therefore, 321

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necessary to soften the structure to allow for shape modification of the bone material. MATERIAL S AN D METHOD S All solvents, solutions and reagents employed in this study were used as received without further purification. Pre-frozen bone specimens from cattle and deer were cut from the condyles of mature beasts using a sharp band saw. Defatting of the bone involved the use of several procedures. In one procedure, the bone samples were effectively autoclaved by boiling for several days in a domestic pressure cooker at ca.l5 psi. Autoclaved bone samples were subsequently immersed in refluxing isopropanol solvent at 82 C in a wide-necked round-bottomed flask for 90 minutes after which excess solvent was removed and the specimens air-dried at ambient temperatures. In another novel procedure, raw bone cubes were placed in ca. 50 mL of water in a beaker and microwaved for 5 minutes in an 800 W domestic microwave oven. The water which contained a significant level of fat was then discarded and the bone cubes subsequently immersed in refluxing isopropanol (82’’C) to extract the remaining fat. Extraction using supercritical CO2 was also trialled. In this procedure, 5 bone cube specimens which had been previously autoclaved as described earlier were placed in a steel reaction vessel and subjected to supercritical CO2 solvent treatment. After 90 minutes of this treatment, the residue collection trap was inspected for fat drawoff and then every 15 minutes after this. After a total two hour treatment time, fat drawoff was not detected and the extraction was judged to be complete. As a comparison, 10 bone cubes not subjected to the previous autoclaving treatment were also subjected to the supercritical CO2 extraction. Fat drawoff from these specimens, in contrast, was still being observed after ca. 3 hours and 15 minutes of supercritical CO2 treatment. Defatted bone specimens destined for clinical use were subsequently subjected to a bleaching process which served to both improve the aesthetic appearance of the bone as well as to decollagenate and thus soften the bone materials in order that they could become cuttable. As with the defatting procedures, several bleaching processes were employed all of which made use of hypochlorite. In the "Cadivar" method of bleaching [1], bone specimens were heated in 1 L of a cloudy solution containing 150 g of Na2C03, 100 g of Ca(0Cl)2 and 150 g of NaOH. Variations on this method involved using lower strength solutions. Other bleaching methods involved the use of commercial strength (1% and 3% NaOCl w/v) bleach solutions. Hydroxyapatite powders were produced by autoclaving partially granulated pieces of bone material and subsequent enzymatic treatment using food grade enzymes such as lipase, nutralase and alcalase to break down extraneous organic matter adhering to the the exterior of the specimens as well as bone marrow. A 10-day dissolution in 5% (v/v) HCl solution ensued. Filtration of scum and undissolved bone material gave a solution which was subsequently treated by stirring in with saturated Ca(0H)2 solution in order to reprecipitate the calcium phosphate. The crude phosphate powder was further purified by redigestion in acid, filtration and subsequent reprecipitation. All bone materials were characterised after the processing steps in the form of ground powders. IR spectra were recorded as KBr disks on a Digilab FTS-40 FTIR spectrometer. ÎP and 13C Solid State NMR spectra were recorded on a Bruker AC200 NMR spectrometer equipped with a solid state probe and MAS using KH2PO4 (for 3lp) and adamantane (for l^c with crosspolarisation) as secondary references. Fat and protein levels in bleached bone specimens were determined using gravimetric petroleum ether soxhlet extraction and Kjedahl analytical procedures. RESULT S AN D DISCUSSIO N Cuttingand Defatting The cutting history of the bone cubes from the condyles was found to be a critical factor

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in achieving efficient defatting. The continued use of a blunt blade led to significant heating of the cutting blade and tended to produce cubes with "sealed" surfaces caused by the conversion of the collagen to gelatin upon contact with the heated metal surface. This effectively blocked pores on the cube thereby preventing efficient defatting. The literature shows that a variety of defatting and deproteination procedures have been used on natural bone [2-6]. Cutting the bones in the frozen state appears to be of importance in past studies [5,6] and was adopted in the present study. However, the use of hydrazine as a deproteination reagent as used by previous workers [2,3] was not deemed suitable for bone destined for eventual clinical use due to concern about hydrazine residues and thus solutions containing hypochlorite were used as a clinically more acceptable alternative. In general, autoclaving of the raw bone samples in water followed by immersion in refluxing isopropanol was judged to be the best method for defatting. Specimens prepared using this procedure and subjected to later deproteination gave the most aesthetically presentable specimens for clinical use. Microwaving of the raw bone specimens was trialled as a novel defatting method due to the fact that microwave treatment will selectively heat the bone matrix immersed in solvent. When initially microwaved in water, bone cubes were observed to heat up in the water which caused streams of liquefied fat to pour out of the bone and collect on the the top of the solution. Removal of water and the subsequent treatment in refluxing isopropanol removed the bulk of the remainder of the fat. Overall, a 58% weight loss was observed in bone cubes treated by this method. Isopropanol was the most effective solvent for fat removal due to its lower polarity compared to solvents such as methanol or ethanol. Microwaving is a highly rapid way of removing bulk fat but an undesired side-effect is excessive heating of the bone in its raw state which can lead to collagen being transformed to gelatin inside the bone thus hindering the penetrability of solvents into the porous structure and hence the defatting and deproteination efficiency. Supercritical CO2 extraction has been used on sheep bone before [6] since CO2 is known to be a good solvent for lipids. Also penetrability of supercritical fluids into porous materials is less of a problem due to elimination of surface tension. The supercritical CO2 extraction studies in the present study demonstrated that some prior treatment of the bones such as autoclaving is essential to cut down on defatting time. The specimens previously subjected to autoclaving were relatively clean in appearance after the supercritical treatment although the "raw" specimens tended to be bloodstreaked and less well-defatted relative to the previously treated samples. Deproteinationand SpectroscopicCharacterisation The cadivar method proved to be an efficient method for efficient bleaching (both inside and outside the cube) and deproteination, but tended to soften the bone cubes excessively to such an extent that they were not clinically acceptable. A problem with bleaching is that deproteination is not uniform so that the outer porous network of bone is weakened more relative to the interior structure. Use of 1% commercial bleach solution was found to be the most acceptable clinically in terms of the strength and aesthetic characteristics. In such samples, the fat and protein content was found to be less than ca. 0.1 g %. Fig. 1(a) is an FTIR spectrum of bovine bone after the microwave(H20)/refluxing isopropanol treatment. The spectrum is typical of bone IR spectra showing features due to collagen, hydroxycarbonate apatite and entrained water, however fat-associated hydrocarbon peaks are extremely weak. Fig. 1(b), in contrast demonstrates the change caused by hypochlorite treatment (using commercial NaOCl solutions). Collagen features (and associated water peaks) are generally absent although the carbonate in the natural bone structure is left intact. In this deviation from the natural bone state, the bone becomes softer and easier to shape to a desired form.

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2000

1

r

1800 1600 1400 1200 Uavenumbers

1 400

Figure 1. FTIR spectra of KBr disks of (a) microwave(H20)/isopropanol-refluxe d bone and (b) microwave(H20)/isopropanol-refluxe d bone subjecte d to NaOC l treatment . In the FTIR spectra of the reprecipitate d hydroxyapatit e from bone, there is predictably an absence of collagen feature s and a much reduced carbonate peak at 1427 cm-1. Although fatassociate d peaks are weak in the reprecipitate d hydroxyapatit e specimens , there is still believe d to be organic matter tenaciousl y held to the surface of the reprecipitate d powders. This was demonstrate d by solid state ^^c NM R spectroscop y which is not often used to characteris e bonederived hydroxyapatit e powders. It was found that in bone or reprecipitate d hydroxyapatit e subjecte d even to hypochlorite treatment , a relatively well-define d complex peak at -30 ppm was observed in 13C NM R spectra which could be due to residual organic component s (e.g. fat and/or collagen/gelatin ) adsorbed strongly to the solid particles at levels not detectabl e by FTIR . Occasionally, weak peaks at 168-17 0 ppm and 185 ppm due to carbonyl groups in collagen and fat respectivel y were also observed . Since the l^c NM R spectrum of this residual organic matter differs from that of pure solid collagen and pure fat derived from bone, this will require further investigatio n by solid state NM R and X-ray photoelectro n spectroscop y to clarify the interpretation . ACKNOWLEDGEMENT S We wish to acknowledg e the New Zealand Foundation for Research, Science and Technology for funding support for this project. We are also grateful to E)r Roger Mederr of Forestry Research Institute of Rotorua for recording of solid state NM R spectra.

REFERENCES 1. 2. 3. 4. 5. 6.

University of Otago Medical School, Dunedin, New Zealand, Private Conmiunication Walters, M.A., Leung, N.C., Blumenthal, R.Z., LeGeros, R.Z., Konsker, K.A., J. Inorg.Biochem,, 1990, 39, 193-200 . Rehman, I., Smith, R., Hench, L.L., Bonfield, W., J. Biomed. Mater,Res.,1995, 29, 1287-1294 . Akazawa, T., Kodaira, K., Phosphorus Res.Bull.,1991,1 , 215-220 . Walsh, W.R., Ohno, M., Guzelsu, N., /. Mater. Sci.Mater. AfeJ.J994 , 5, 72-79. Frayssinet, P., Asimus, E., Autefage, A., Fages, J., J. Mater. Sci.Mater.Med.,1995, 6, 473-478 .


CATASTROPHI C WEA R O F META L BAL L O F BIPOLA R HI P PROSTHESI S AFTE R FRACTUR E O F ALUMIN A CERAMI C SCREW S USE D FOR ACETABULA R BON E GRAF T H. Ohashi, Y. Yutani, A. Kobayashi, Y. Kadoya and Y. Yamano Department of Orthopaedic Surgery, Osaka City University Medical School, Asahimachi, Abenoku, Osaka 545, Japan ABSTRAC T Adverse effect of alumina ceramic was investigated in a patient with a rapid progress of prosthetic loosening after catastrophic wear of metal ball following a fracture of alumina screws used for acetabular bone graft. Alumina ceramic fragments were detected on the bearing surface of polyethylene insert. The fragments were considered to rasp the surface of metal ball producing abundant sub-micron metal particles. The levels of II-la, II-1P, 11-6 and TNF-a in the joint fluid were high. The progress of prosthetic loosening was supposed to be accelerated by production of the cytokines. Since alumina ceramics, especially these of screw shaped, are brittle, we concluded that the use of alumina ceramic screws in prosthetic replacement was contraindication. KEYWORDS : alumina ceramic, metallosis, total hip replacement, wear, metal particles, loosening INTRODUCTIO N Alumina ceramic fragments exsisted in articulating interface of joint prostheses can cause typical third-body abrasive wear. Several cases of severe metallosis have been reported after fracture of alumina ball [1] and alumina screws [23] In these cases, prosthetic loosening was observed, however little has been discussed about the mechanism of loosening. Recently, small polyethylene (PE) particulates are reported to play a great role to induce periprosthetic osteolysis by activating macrophages [4,5]. In case of metallosis, the mechanism of loosening is not well recognized. We experienced a rapid progress of loosening after catastrophic wear of metal ball of a bipolar hip prosthesis with a fracture of alumina screws. The aim of this study was to reveal the mechanism of the catastrophic wear and to investigate the relationship between a fracture of the alumina screws and prosthetic loosening. This adverse effect of alumina ceramics may warn its clinical applications in certain conditions. MATERIAL S AN D METHOD S A 47-year-old woman imderwent a right bipolar hip hemiarthroplasty (Bateman UPF, Co-Cr alloy) due to bilateral coxarthrosis. Acetabular dysplasia was supplemented with a bone graft fixed by two alumina ceramic screws (Sapphire Screw, monocrystal alumina ceramic) (Fig. 1). Her postoperative course was uneventful, howener proximal migration of the outer head was observed on serial radiographs. Nine years postoperatively, radiographs revealed a fracture of alumina screws due to impingement with the migrated outer head. Five months later, she felt severe coxalgia (Fig. 2) and a revision arthroplasty was performed. 325

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F i g . 1 Radiograph taken after the first operation.

F i g . 2 Radiograph taken before revision surgery.

At surgery, black joint fluid was collected, and the capsule and the periprosthetic tissue were diffusely stained with a black material. The metal ball wore out like a rugby ball (Fig. 3) and the femoral component was loosened. There was no evidence of infection. Geometry of metal ball and bearing insert was analysed by a coordinate measuring machine (BHN305, Mitutoyo Co.). Surfaces of retrieved metal ball and outer head were examined using scanning electron microscopy (SEM). Debris sticked to the bearing surface of PE insert was analysed by energy-dispersive analysis of X-rays (EDAX). Metal particles were extracted from the joint fluid by tissue digestion [6]. The retrieved metal particles in SEM photographs were measured by a computerized image analyzer and the size was estabhshed using the equivalent circle diameter (ECD). The levels of I l - l a , II-ip, 11-6, TNF-a in the joint fluid were measured by radioimmuniassay. RESULT S Volmetric wear of the metal ball was 729mm3. and that of the PE insert was 236mm3. SEM of the metal ball represemted many sharp scratches (Fig. 4). SEM of the bearing surface of PE insert represented roughened surface with two kinds of debris. EDAX revealed that they were

F i g . 3 Severely worn metal ball.

F i g . 4 SEM of the surface of metal ball.

Catastrophic Wear of Metal Ball of Bipolar Hip Prosthesis:H. Ohashi et al. 327 >K4 .

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metal and alumina (Fig. 5, 6). Mean BCD of the metal particles was 0.99–0.86 ;^m, and the size distribution was shown in Fig. 7. Histologically, metal debris as well as PE debris were diffusely observed in the periprosthetic tissue (Fig. 8). The levels of I l - l a , II-ip, 11-6 and TNF-a in the joint fluid were 42.5 ng/ml, 90.5 ng/ml, 3040.3 pg/ml and 73.0 pg/ml, respectively. DISCUSSIO N AN D CONCLUSION S Alumina ceramic has good biocompatibility, while alumina tends to break to small fragments due to their brittleness. Hardness of alumina ceramic fragments is greater than that of Co-Cr alloy, thus alumina ceramic fragments can rasp the surface of metal ball. From the observation of retrieved prosthesis, the mechanism of the catastrophic wear in this case is considered to be an abrasive wear due to alumina fragments sticked to the bearing surface of PE insert. Recently, sub-micron particulates are considered to play an important role in periprosthetic osteolysis [5,7], and bone-implant interface around failed joint prostheses was reported to contain high level of cytokines especially in regions with radiographic evidence of bone loss [8]. Experimentally, Co-Cr particles were reported to induce proliferation of macrophages [9], and stimulate cytokine production [10]. In this case, most metal particles were sub-micron in size and high levels of Il-la, II-1 (3,11-6 and TNF-a, that are associated with bone resorption, were detected in the joint fluid. These results indicated that the fracture of alumina ceramic screws brought about catastrophic wear of metal ball. The abundant sub-micron metal particles stimulated the production of cytokines that supposedly accelerated the progress of prosthetic loosening. From this point of view, we concluded that the use of alumina ceramic screws was contraindication in prosthetic replacement.

REFERENCES 1. Kempf, I. and Semlitsh, M., Arch. Orthop.TraumaSurg., 1990, 109, 284-287. 2. Matsuda, Y , Yamanuro, T., Kasai, R., Matsusue, Y. and Okumura, H., J. Arthroplasty,1992, 7S, 439-445. 3. Watanabe, M., Okumura, H., Kihara, Y. and Shibata, T., Arch. Orthop.TraumaSurg., 1993, 112,296-298. 4. Chiba, J., Rubash, H., Kim, K.J. and Iwaki, Y., Clin. Orthop.,1994,300,304-312. 5. Shanbhag, A.S., Jacobs, J.J., Giant, T.T., Gilbert, J.L., Black, J. and Galante, J.O., J. Bone Joint Surg., 1994,76-B, 60-67. 6. Yamac, T., Kobayashi, A., Bonfield, W., Kadoya, Y., Freeman, M A R . , Scott, G. and Revell, RA., Transactionsof Fifth WorldBiomaterialsCongress,1996, 861. 7. McKellop, H.A., Campbell, P., Park, S.H., Schmalzried, T.P., Grigoris, P., Amstutz, H.C. and Sarmiento, A., Clin. Orthop.,1995,311, 3-20. 8. Chiba, J., Schwendeman, L.J., Booth, RE., Crossett, L.S. and Rubash, HE., Clin. Orthop., 1994,299,114-124. 9. Howie, D.W. and Vernon-Roberts, B., Clin. Orthop.,1988,232, 244-254. 10. Dowd, J.E., Schwendeman, L.J., Macaulay, W., Doyle, J.S., Shanbhag, A S . , Wilson, S., Herdon, J.H., Rubash, HE., Clin. Orthop.,1995,319, 106-121.


ANTIBACTERIA L PROPERT Y OF Ag-DOPE D CALCIU M PHOSPHAT E COMPOUND-CELLULOS E COMPOSITE S K. Okada\ Y. Yokogawa^ T. Kameyama^ K. Kato^ Y. Kawamoto^ K. Nishizawa^ F. Nagata^ and M. Okuyama^ ^R & D Center, NGK Spark Plug Co., Ltd., 2808 Iwasaki, Komaki, Aichi 485 Japan ^Bioceramic Laboratory, National Industrial Research Institute of Nagoya, 1-1 Hirate-cho, Kita-ku, Nagoya 462 Japan ^Laboratory of Bioorganic Chemistry, National Industrial Research Institute of Nagoya, 1-1 Hirate-cho, Kita-ku, Nagoya 462 Japan ABSTRAC T Ag-doped calcium phosphate compound (CP)-cellulose fiber composites were prepared by immersing CP-cellulose fiber composites into AgNGs aqueous solution after depositing CP on the cellulose fibers by soaking Ca(0H)2-treated phosphorylated cellulose into 1.5xSBF(Stimulated Body Fluid). The composite obtained became dark-brown from white gradually as Ag"*^ ion concentration in the AgNOa aqueous solution increased. The decrease of Ag* ion concentration in the AgNOa aqueous solution was also observed by ICP analysis and the broad XRD peak due to CP was shifted slightly for Ag-doped CP-cellulose fiber composites. It appeared from these that Ag was doped into the CP lattice. The antibacterial property of the composites was investigated by using Bacillus subtilis. As the amount of doped-Ag increased, the growth of Bacillus subtilis was inhibited. INTRODUCTIO N Hydroxyapatite (Caio(P04)6(OH)2); HAp) has been used as bone substitutes and dental implants because of its similar structure as the mineral phase in bone and teeth and has a high affinity for living bone. HAp has been also used as an adsorbent for high performance liquid chromatography column because of its adsorptive properties for virus, bacteria, and protein[l]. On the other hand, metals such as silver, copper, and zinc were well-known to have antibacterial property and antibacterial products have been being fabricating by mixing these metals-doped powders such as zeolite and HAp with fibers or resins. In the previous paper, calcium phosphate compound (CP) growth on cellulose fibers phosphorylated in 1.5xSBF was studied[2]. CP-cellulose fiber composites doped with these metals in the CP layer are expected to be useful as a virus and bacteria adsorptive filter with antibacterial property. In this present study, Ag-doping was attempted into CP layer of CP-cellulose fiber composites, and their characteristic and antibacterial property were evaluated. 329

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EXPERIMENTA L Deposition of Calcium Phosphate Compound Phosphorylation of cellulose fibers (Cotton Ciegalfi;100 % cotton) was carried out following the method described in an earlier report[2]. 16 pieces of cellulose fibers (ca. 5.5 g) were placed into a round-bottomed flask with 40 g of urea and 250 mL of DMF, and heated up to 120*0 with mechanical stirring under N2 atmosphere, and then approximately 32.0 g of 85 % H3PO4 in 100 mL of DMF was added into the solution. The temperature of the reaction mixture was increased to 155*0 and refluxed for 1 h. The reaction mixture was then left to cool under N2 gas flow, and the cellulose fibers were filtrated and washed thoroughly in water to rinse out excess phosphoric acid. The phosphorylated cellulose fibers were soaked without stirring in a saturated solution of ca. 500mL of Ca(0H)2 (pH = 12.5) in 1000 mL closed plastic bottle for 8 days. The Ca(0H)2 solution was renewed every 4 days. After soaking, the fibers were washed with pure water, and dried at 60*C under vacuum. 0.1 g of Ca(0H)2-treated phosphorylated cellulose fibers were The 1.5xSBF was prepared in immersed into the 200 mL of 1.5xSBF for 8 days at 3 6 . 5 0 . the same manner of an earlier publication[3]. The 1.5xSBF solution was basically prepared by dissolution of NaCl, KCl, CaCh, MgCh, NaHCOa, NaS04, and K2HPO4 in pure water with buffering agents "TRIS" ((CH20H)3CNH2) and HCl to keep the solution pH of 7.2 - 7.3 during soaking experiments. After soaking, the cellulose fibers were washed with pure water and dried at 6OO under vacuum. Ag-doping and Characterizatio n Ag-doping were carried out with two methods as follows. For method A, the CPcellulose fiber composites prepared were soaked into AgNOa aqueous solutions at ambient temperature for 1 day, where the Ag amounts were 0.5, 1, and 3 mol% against to Ca amount in the CP deposited. After soaking, the cellulose fiber composites were washed with pure water and then dried. For method B, the Ca(0H)2-treated phosphorylated cellulose fibers were soaked into 1.5xSBF containing Ag"^ ion at 3 6 . 5 0 for 8 days, where the amounts of Ag"*^ ion were 1, 5, 10, and 30mol% against to Ca^"*^ ion in 1.5xSBF prepared by using nitrate salts instead of chloride salts. The microstructure, the amount of Ag doped and crystalline phase were characterized with scanning electron microscopy (SEM), energy dispersive X-ray (EDX) and X-ray diffraction (XRD) analyses. The Ag concentration change in AgNOa solutions were studied by ICP analyzer. Evaluation of Antibacterial Property Antibacterial property was evaluated by following method. Yeast extract, polypepton, MgS04 and agar were added into O.IM phosphate buffer solution and dissolved by boiling. After sterilizing the solution with an autoclave, each 10ml of the solution were pipetted into sterilized dishes. The Ag-doped CP-cellulose fiber composites were put on the culture medium which 10^ pieces of Bacillus subtilis was set on, and then

AntibacterialProperty of Ag-Doped CP Compound-CelluloseComposites:K. Okada et al.

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cultivated at 37 C for 24 h. After cultivation, the growth of Batillus subtilis around the sample and under the sample was evaluated. RESUL T AN D DISCUSSIO N The composites prepared by method A changed from white to dark-brown gradually as Ag^ ion concentration in the AgNOa aqueous solution increased. However, Ag amount in/on the composites was too small to be detected by EDX. The decrease of Ag"*" ion concentration in the AgNOa aqueous solution was observed by ICP analysis, and also the broad peak due to CP was shifted slightly for Ag-doped CP-cellulose fiber composites. It appeared from these that Ag was doped into CP lattice. Major difference was not observed in the microstructure such as morphology and size of primary grain. On the other hand, the weight increase of the composites prepared by method B were observed, but the amount of the compound deposited on the fiber decreased as the amount of Ag"*^ ion in the solution increased. Especially no deposition of CP was observed when the solution including 30mol% Ag"*^ ion was used. The color change to dark-brown and the shift of XRD peak due to CP were not observed differently from method A. ICP analysis didn’t show any changes in Ag concentration in the solution. It was considered to be difficult to prepare Ag-doped CP-cellulose composite by using l.SxSBF including Ag^ ion. Antibacterial properties were evaluated for Ag-doped CP-cellulose fiber composites prepared by method A, CP-cellulose fiber composite, and cellulose itself. In Figure 1, the inhibition of the growth of Bacillus subtilis was indicated for Agdoped CP-cellulose fiber composite, while it was not observed for cellulose and CPcellulose fiber composite. Also as the amount of Ag doped into CP increased, the inhibition of the growth of Bacillus subtilis was remarkable around the sample as well as under the sample.

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Tinhibited area l Figure 1 Photographs of growth of Bacillus subtilis around the samples (a) Cellulose, (b) Calcium Phosphate Compound-Cellulose Composite, (c) Agdoped Calcium Phosphate CompoundCellulose Composite SUMMAR Y Ag could be doped into calcium phosphate compound (CP) deposited on cellulose fibers by immersing CP-cellulose fiber composites into AgNOa solution and Ag-doped CPcellulose fiber composites had antibacterial property for Bacillus subtilis qualitatively. It is believed that these composites can be applied to the filter which have antibacterial property.

REFERENCES l.Tsuru S., Shinomiya N., Katsura Y., Uwabe Y., Noritake M. and Rokutanda M., BioMedical Materials and Engineering, 1991, 1, 143-147 2.Mucalo M.R., Yokogawa Y., Toriyama M., Suzuki., Kawamoto Y. and Nishizawa K., Journal Material Science, Material in Medicine, 1995, 6, 658-669 3.Li P., Otsuki C , Kokubo T., Nakanishi K., Soga N., Nakamura K. and Yamamuro T., Journal Material Science, Material in Medicine, 1993, 4, 127-131


WEA R BEHAVIOU R OF POLYETHYLEN E CUP AGAINS T 28mm ALUMIN A BAL L IN TOTA L HI P PROSTHESE S H. Oonishi. N. Murata. S. Kushitani, 8. Wakitani. K. Imoto. Y. Iwaki. and N. Kin. Department of Orthopaedic Surgery. Artificial Joint Section and Biomaterial Research Laboratory. Osaka-Minami National Hospital. 2 - 1 . Kidohigashimachi. Kawachinagano-Shi. Osaka. 586. JAPAN

ABSTRACT The wear rate of the U. H. M. W. polyethelene cups in combination with 28 mm alumina femoral head was measured on the radiographs whithout any complications and on the retrieved cups due to slight loosening of the p r o s › theses or due to late infections between bone and components. In both cases, the thicker the polyethylene cups, the lower the wear rate. The average wear rate of the cups of 7 and 8 mm thickness was about twice of that of 10 and 11 mm. From these results, the thickness of the polyethylene cups have to be used more than 11 mm. KEYWORD S wear of polyethylene cup. polyethylene cup thickness, total hip p r o s t h e › sis, alumina head INTRODUCTIO N We reported previously that the wear rate of the cup on 28 mm metal head(T-28. stainless steel head ball) was 2.5 times of that on 28 mm alumina head (Bioceram) [1]. The objective of our study was to find the relationship between the wear and the polyethylene cup thickness of cemented alumina ball total hip prostheses from the radiographs and the retrieved cups. We reproted previously the effect of the polethylene cup thickness of 7 to 9 mm to wear on the radiographs [1]. In this sutdy the polyehtylene cup thickness of 7 to 11 mm was compared on the radiographs and on the retrieved cups. CU P WEA R ON THE RADIOGRAPH S 1) Materials 111 joints in 102 cases were considered suitable for inclusion in this study. 14 joints in 13 cases were in male and 97 joints in 89 cases were in 333

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female. All cases were secondary osteoarthritis due to dysplasic acetabulum. 93 cases were unilatereal and 9 cases were bilatereal. 2) Methods The X - r a y were taken of A - P views of both hips in standing position (weightbearing). We used the decrease in thickness of the UHMWPE cup as our estimate of wear. Observation periods were from 1 month after s u r › gery and the longest year after surgery. For measurement on the radiographs, we used backlit-type digitizers with 20urn resolution, 5 x magnification view› ing loupes, and specially designed angle scales (0. 2 mm graduation). Data collation and analysis wear via computer. In our previous studies of the measurement of the wear of the polyethylene cups on the radiographs, we found the tendency that the thicker the cups, the lower the wear of the cups. The similar results were reported in hip simulator tests by Saikko [2]. T h e r e › fore, in this study the relationships between the cup thickness and the wear were investigated. Volumetric wear rate was calculated from linear wear rate using a system developed by J. Michael Cabo et al [3]. 3) Results The linear wear rate and the volumetric wear rate of each cases were shown in the figure 1 and 2 . The wear rate of each cases were scattered very widely. The average wear rate of the cups of 7 and 8 mm thickness was about twice of that of 10 and 11 mm. On the whole, the thicker the cup, the lower the wear rate. r

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Wear Behaviour of PE Cup Against 28 mmAlumina Ball In Total Hip Prostheses:H. Oonishi et al. 335

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CUP WEAR ON THE RETRIEVED CUPS 1) Materials T h e d e c r e m e n t of t h e cup t h i c k n e s s , including w e a r and c r e e p d e f o r m i t y , of Bioceram p r o s t h e s e s w a s m e a s u r e d on t h e r e t r i e v e d prostheses. The retrieved prostheses were due to slight loosening of the stem, cup or both, or due to late infection between bone and components. Prostheses damaged by other than wear by femoral head were excluded from this study. 2) Methods Generally, the inner surface of the retrieved cup has two spherical s u r › faces. The distance between two centers was defined as t h e length of t h e femoral head movement. In this case, as the initial wear, which was extremely higher than the steady steady state wear, could not be excluded, the wear rates included the initial wear in the steady state wear rate. 3) Results Relationships between linear and volumetric wear rates and cup t h i c k › ness were shown on the figure 3. The wear rates of the cups of 7 mm and 8 mm thickness were almost the same. The wear rates of the 7 and 8 mm cup thickness was twice of that of the 11 mm cup thickness. The thicker the cup. the lower the wear rate. DISCUSSIO N In the case on radiographs, if one case of the 44 mm in thickness and two cases of 50 mm in thickness, which showed extremely higher wear rates, were excluded from the population, the average line on the on the graph will get nearer to slow sloping or straight line. In the case on retrieved cups, the

336

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Fig. 3 ; Linear and volumetric wear rate of polyethylene cups on the retrieved prostheses. Relationships between wear rates and cup thickness. average line on the graphs showed a slow sloping curve. In both studies on wear measurement on the radiographs and on retrieved prostheses, similar t e n d › ency and reasonable results were obtained in relationships between cup t h i c k › ness and wear rate of the cups. As a result, on the whole, the thicker the polyethylene cups, the lower the wear rate. The average wear rate of the cups of 7 and 8 mm thickness was about twice of that of 10 and 11 mm. The same tendency was reported in hip simulator test using water lubricant by Saikko [2]. However, the wear rate on the retrieved cups was higher by 50% than that on the radiographs. Because, in the retrieved cases, as the prostheses were not removed from the patients without any complications, but removed from the patients suffering from loosening of the components or late infections, higher wear rate was supposed to be found. Moreover, in the case on the r a › diographs, the initial wear, which was extremely higher than the steady state wear, was excluded However, in the case on the retrieved cups, as the initial wear was included, the wear rate in the case on retrieved cups became higher than that in the case on the radiographs. From these results, the thickness of the cups must be used more than 11 mm.

REFERENCES 1) OONISHI, K , TAKAYAMA, Y.. CLARKE I. C. and JUNG H.; J. of LongTerm Effects of Medical Implants. 1992. 2(1). 3 7 - 4 9 . 2) SAIKKO. v.. ; Acta Orthop. Scand. 1995. 66. 501-506. 3) CABO. J. M, ; J. of Bone and Joint Surgery. 1993, 7 5 - B ( 2 ) , 2 5 4 - 2 6


IN VITR O CEL L BEHAVIO R OF OSTEOBLAST S ON PYROS T BON E SUBSTITUT E J.S. Sun^ , F.H. L i n ’ , Y.H. Tsuang \ Y.S. Hang \ C.Y. Hong \ and H.C. Liu’ ^ Department of Orthopedic Surgery, National Taiwan University Hospital, and ’ Center of Biomedical engineering. College of Medicine, National Taiwan University, Taipei, Taiwan.

ABSTRACT

We have, elucidated the in vitro cell behavior of osteoblasts on Pyrost bone substitute. Using primary culture of rat osteoblasts, the changes in cell morphology on the surface of Pyrost bone substitute were studied. At 1 hour, 3 hours and days 1, 3, 7 after layering, the cell behavior was observed with SEM. The processes of trypsinized osteoblasts adhesion and spreading on Pyrost bone substitute consisted of: 1). contact of rounded osteoblasts with the Pyrost substrate; 2). attachment of osteoblasts at point of contact; 3). centriftigal growth of filopodia; 4). flattening and spreading of the osteoblasts on the Pyrost substrate; 5). division and growth of osteoblasts; 6). suspension of the osteoblasts across the pores by their processes. This result demonstrated that Pyrost can form a physico-chemical bond with osteoblasts. The Pyrosts bone substitute can support both attachment and proliferation of osteoblasts. KEYWORDS : pyrost bone, osteoblasts, adhesion, spreading.

INTRODUCTION

Approaches to bone regeneration for the treatment of various clinical conditions, such as fi-actures with bone loss, bone infections or bone tumors, involve the use of autogenous grafts or allografts [1]. Autogenous cancellous bone is the most effective bone graft material to date, but it also has drawbacks, including donor site morbidity and limited availability, especially in children [2]. The advantage of allografts over autografts lies in better availability of supply and their ability to be used for reconstructing large bone defects [1]. The major disadvantages of allogeneic bone include disease transmission and the graft’s tendency to elicit an immune response that can lead to high failure rates [3]. Pyrost bone substitute has been shown to be a promising orthopedic biomaterial. When used as a bone graft substitute, bony ingrowth into the implants without any adverse reaction can be demonstrated [4]. When implanted into bone, Pyrost can form a physico-chemical bond with bone tissue. However, little is known about the mechanisms responsible for the osteogenesis that occurs between bone and Pyrost bone substitute. The in vivo as well as in vitro bone formation are closely associated with the behavior of the cells. The formation and deposition of bone directly on to the implant require a surface that is not only non-toxic but also allows or favors the cell behavior [5]. 337

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We therefore initiated studies of the morphology of osteoblasts to test the in vitro cell behavior on the Pyrost bone substitute. MATERIALS AND METHODS

Sequential digestion of newborn Wistar-rat calavria was performed by using a modification of the methods described by Wong and Cohn [6]. Briefly, the dissected calvaria were sequentially digested with collagenase (180 U/ml, Sigma Co.) in solution A \yith EDTA. The cells released after last two treatments were harvested by centrifugation and resuspended in culture medium. The Pyrostfi (Oscal AG, Swiss) bones (10 x 5.0 x 5.0 mm) were placed in sterile dishes. Confluent rat osteoblast cultures were seeded into each well on top of the implant materials at a density of 3 x 10’* cells/cm^ The culture media used was DMEM supplemented with 10% fetal calf serum (Gibco, UK) and antibiotics (penicillin G sodium 100 units/ml -streptomycin 100 mg/ml, Gibco, UK). The dishes were incubated at 37 C in an atmosphere supplemented with 5% CO2 and fed with complete changes of medium twice a week. The day of plating was considered as the zero day of culture. The test samples were removed from wells at 1 hour, 3 hours and days 1,3,7, fixed in 3% formaldehyde in 0.1 M PBS buffer (pH 7.4). For electron microscopic examination, the Pyrost bone substitute blocks were fixed, dehydrated and critical dried. Specimens were sputter-coated with gold and examined by scanning electron microscope. RESULTS AND DISCUSSION

After layering on the Pyrost, the scattered round-shaped osteoblasts settled on the substratum with the proteinaceous sheets within 3 hours. One day after layering, flattening of some osteoblasts were visible. After 3 days in culture, the cells exhibited close contact with each other via filopodial processes. The surface of Pyrost was coated by an almost complete layer of osteoblasts by the day 7. The cells initially repopulate the Pyrost bone substitute by settling out of suspension, attach to the available surfaces provided, and then give rise to the final populations by mitotic expansion. In vivo, this condition is mandatory for osteogenesis to occur in an implanted material without interposition of fibrous or granulation tissue [5]. In this series, after trypsinization, the cells appeared spherical to ovoid in shape. The population consisted of cells with smooth, rounded surface, and surface that possessed numerous ’bleb’-like vesicular protuberances within 1 hours after layering (Fig. 1). Some cells showed smooth surface without microvilli or blebs. The numerous foldings and blebs on the surface of harvested cells are to accommodate the excess surface membrane as the cells round up from the flattened state in response to trypsin treatment [7]. The major events in the process of adhesion and spreading of these cells seem to be attachment of the cell to the substratum, radial growth of filopodia, cytoplasmic webbing and the resultant flattening of the cell. When cells layered on Pyrost bone substitute after 3 hours, the cells adhered very firmly to the surface. This was affected by microvilli-like cell processes. Their growth occurred only at the point of contact with the Pyrost (Fig. 2). It is likely that the microvilli-like projections were formed all over the cell surface but were later withdrawn except at point of contact with the substratum. The first step in cell spreading constitutes the centrifugal growth of microvilli-like processes that elongated into filopodia (Fig. 3). It appears that these spherical tips of filopodia may play a direct role as specialized structure of attachment to the substratum [7].

In Vitro Cell Behavior of Osteoblastson Pyrost Bone Substitute:J-S. Sun et al.

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I’ig. 1 S E M exaJTiination of trypsniharvested osteoblasts fixed vvilbin 1 hour after layering. (Bar: 2,6 |Lim)Fig. 2 SHM examination oftrypsijiharveslcd osteoblasts fixed at 3 hours after layering. Spherical cells with inconspicuous microvilli-]ike projections (Bar: 1.9 îrn). Fig. 3 SUM examination of trypsinharvestcd osteoblasts fixed at 3 hours after layering. The contact area with long filopods- (Bar: 2.S ^ m ) .

Figs. 4-7 SEM examination of trypsin-harvested osteoblasts. Fig. 4 One day after layering: Flattened cell with cytoplasmic webbing (Bar: 4.4 |Lim). Fig. 5 Three day after layering. Cells with the long filopods (Bar: 5.0 fim). Fig. 6 Three day after layering. The division and growth of osteoblasts (Bar: 3.9 |Ltm) Fig. 7 Seven day after layering. Cells lay densely on Pyrost; and cells spanned the pores apparently by first expanding explorative filopods across the macropore. (Bar: 13.0 jam).

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Close contacts made by the advancing filopodia is followed by extensive spreading as a thin web between the filopodia (Fig. 4). By the first day after layering, cells have flattened considerably. It is reasonable to conclude that the strength of adhesion of a cell settling on the Pyrost bone substitute would increase progressively as the area of its contact increases (Fig. 5). Some of the blebs still persisted in the surface of the flat cell. The growth of filopodia followed immediately by cytoplasmicwebbing seems to be the pattern of origin of the ruffling membrane. The origin of ruffling membranes may indicate the final stage of spreading and the beginning of the cell movement [7]. Following the cell attachment, division and growth of cells occurred at 3 days after layering (Fig. 6). During mitosis a reversal of these events occurs by cytoplasmic de-webbing; the cell becomes spherical. After mitosis, the two daughter cells become flat again by cytoplasmic webbing. At the 7ih day, proliferating cells lay flat and in close contact with the Pyrost surface (Fig. 7). On the other hand, cells appeared suspended across the pores by their processes (Fig. 7). As a result, the cells were not in close contact with the underlying substratum. The formation and deposition of osteoblasts directly on to the Pyrost surface imply that it is not only non-toxic but also allows or favors the osteoblasts behavior. SUMMARY

Surface reactivity may be a key factor in determining the morphological and functional responses observed during the osteoblast-substrate interactions. Pyrosts bone substitute is considered osteo-compatible [4]. This experiment has defined that Pyrosts bone substitute is not only support osteoblasts attachment but it also allows proliferation of the cells. It should be emphasized that these different stages are not discretely separable but are different phases of a contiguous process. This is not a synchronous cell population so variation in the duration of these phases exists and the degree of overlapping of these events is observed. Complete interpretations of these events, however, require further investigation of both morphological and functional responses of osteoblasts to Pyrost bone substitute, in particular with human osteoblasts. ACKNOWLEDGEMENT

S

The authors sincerely appreciate the National Science Council (ROC) for their financial support to accomplish the research. REFERENCES 1. Mankin, H.J., Gebhardt, M.C., and Tomford, W.W., Orthop. Clin. North Am., 1987, 18, 275289. 2. Begley, C.T., Doherty, M.J., Hankey, D.P., and Wilson, D.J., Bone, 1993, 14, 661-666. 3. Bos, G.D., Goldberg, V.M., Zika, J.M., Heiple, K.G., and Powell, A.E., J. Bone Joint Surg., 1983, 65A, 239-246. 4. Katthagen, B,D. 1986 Bone regeneration with bone substitutes: An animal study. Springer, Berlin Heidelberg New York, 29-50. 5. Bagambisa, F.B., and Joos, U., Biomaterials, 1990, 11, 50-56. 6. Boonekamp, P.M., Kekkelman, J.W., Hamilton, J.W., Cohn, D.V., Jilka, R.L., Proc. Kon. Acad. Wet. B., 1984, 87,: 371-384. 7. Rajaraman, R., Rounds, D.E., Yen, S.P.S., Rembaum, A., Exptl. Cell Res., 1974, 88, 327-339.


TH E EFFICAC Y OF HYDROXYAPATITE-TRICALCIU M PHOSPHAT E FILLE R FO R BON E DEFECT S ASSOCIATE D WIT H HUMERA L PSEUDOARTHROSI S : COMPARISO N WIT H AUTOGENOU S ILIA C BON E GRAFT S Katsuji Suzuki and Mitsuko Yamada. Department of Orthopaedic Surgery, Fujita Health University School of Medicine. 1-98, Dengakugakubo, Kutsukake, Toyoake, Aichi, 470-11, Japan. ABSTRAC T The efficacy of a hydroxyapatite-tricalcium phosphate ( HAP-TCP ) AIICT for bone defects associated with humeral pseudoarthrosis was studied by comparing 7 patients treated with HAP-TCP and 7 who received iliac autografts. There WCTC no significant differences in age, additional injuries, and the non-union period between the HAP-TCP group and the iliac autograft groups. Tha-e were also no significant differences in the postoperative bone union time and range of motion ( ROM ) recovery time. KE Y WORD S hydroxyapatite-tricalcium phosphate humeral pseudoarthrosis iliac autograft bone union range of motion INTRODUCTIO N In patients with humeral pseudoarthrosis, it is necessary to resect and freshen the sclerotic bone and fibrous tissue, as well as filling bone defects. Iliac autografts are often used, but present problems due to the limited amount of bone that can be harvested and the occurrence of symptoms at the donor site. We compared the effectiveness of hydroxyapatite-tricalcium phosphate ( HAP-TCP ) and iliac autografts for achieving bone union in patients with humeral pseudoarthrosis. SUBJECT S Fourteen patients with humâl pseudoarthrosis and bone defects requiring filling underwent surgery at this department between 1987 and 1995 ( 9 males and 5 females ; average age : 35.1 years; range : 7 - 75 years ). HAP-TCP group HAP-TCP was used in 7 patients ( 5 males and2 females; average age : 35.4 years ; range : 7-69 years ). The avaage period between injury and surgay for psedoarthrosis was 70.3 + 26.6 ( M – S E ) months and the range was 12 to 240 months. Two of them had shaft pseudoarthrosis, 2 had lateral condylar non-union, 1 had comminuted concfylar non-union, 1 had neck non-union, and 1 had supracondylar non-union. In addition, there was tardy ulnar nerve palsy in 1 patient, radial nerve palsy in 1, axillary nerve palsy in 1, and osteoporosis in 1. 341

342


Iliac autograft group Iliac autografts were used in 7 patients (4 males and 3 females ; average age :34.9 years; range 11 - 75 years). The average period until surgery for non-union was 108.0–59.5 ( M – S E ) months and the range was 12 to 456 months. Three of these patients had lateral condylar non-union, 2 had shaft non-union, 1 had medial condylar non-union, and 1 had supracondylar non-union. In addition, there was tardy ulnar nerve palsy in 2 patients and osteoporosis in 1. METHOD S In both groups, we measured the time until bone union was apparent on X-ray (bone union time) and the time required for the range of elbow joint motion to recover to greater than 80% of that on the healthy side ( ROM recovery time ).

Figure 1. Case 1. A 30-year-old man from the HAP-TCP group. 1-A: Pseudoarthrosis of the shaft of the right humerus 34 months after injury. Ender nailing was done twice, but pseudoarthrosis of the humeral shaft was developed 1-B: Four weeks after surgery. The bone defect at the site of pseudoarthrosis was filled with an HAP-TCP block and granules. 1-C: One year after surgery, bone union is good.

Efficacy of Hydroxyapatite-TricalciumPhosphateFiller for Bone Defects: K. Suzuki and M. Yamada 343

RESULT S The bone union time was 23.4–4.6 ( M – S E ) weeks in the HAP-TCP group and 19.4–3.6 weeks in the iliac autograft group (P=N.S.; Wilcoxon test). The ROM recovery time was 17.9–4.3 ( M – SE ) weeks in the HAP-TCP group and 13.9 – 1.4 weeks in the iliac autograft group (P=N.S.; Wilcoxon test). Postoperative infection did not occur in either group. Pain at the site of bone removal occurred in 3 patients from the iliac autograft group, while thê were no complications in the HAPTCP group. CAS E REPORT S Case 1 was a 30-year-old man ( HAP-TCP group) with pseudoarthrosis of the right humeral shaft. He suffered a right humeral shaft fracture in a traffic accident and was operated on twice with Ender nails at another hospital. However, bone union was not obtained afto" 34 months. At our department, the Ender nails were removed and the sclerotic bone and fibrous tissue at the pseudoarthrosis were resected. Then an HAP-TCP block and granules were used to fill the bone defect and firm internal fixation was achieved with a titanium plate and screws. The bone union time was 24 weeks and the ROM recovery time was 6 weeks.

Figure 2. Case 2. A 34-year-old man from the iliac autograft group. 2-A: Pseudoarthrosis of the shaft of the right humerus 30 months afta* injury. 2-B: Immediately aftCT surgery. The bone defect at the pseudoarthrosis site was filled with an iliac bone block and cancellous bone chips. The bone defect at the site of intCTlocking nail removal was filled with HAP-TCP granules. 2-C: Twelve weeks after surgery, bone union is good.

344


Case 2 was a 34-year-old man (iliac autograft group ) with pseudoarthrosis of the right humeral shaft. He suffered a right humeral shaft fracture after falling from a height, and was treated with an interlocking nail at another hospital. However, bone union was not obtained after 30 months. At our department, the pseudoarthrosis was resected and freshened, and the interlocking nail and a screw were removed. An iliac bone block and cancellous bone chips woe used to fill the bone defect, and firm internal fixation was achieved with a titanium plate and screws. The bone union time was 12 weeks and the ROM recovery time was 8 weeks. DISCUSSIO N HAP-TCP filler is a composite of hydroxy apatite (HAP) and tricalcium phosphate (TCP) which shows excellent biocompatibility and bone conductivity. TCP also fuses strongly with the surrounding bone because it acts as a bone substitute. Suzuki et al.,[l] packed traumatic bone defects of tibia with HAP-TCP fill^ and compared its efficacy with that of autogenous bone grafts. The weight-bearing recovay time was significantly shorter in the HAP-TCP group than in the autogenous bone group. Suzuki etal.,[2] also filled traumatic bone defects of the distal radius with the HAP-TCP filler and compared its efficacy with that of iliac autografts. Although osteoporosis was significantly more conmion in the HAP-TCP group, there was no significant dififoience in the ROM recovery time and grip powCT recovery time. In addition, the 1-year postopâtive rado-ulnar distance was significantly greater in the HAP-TCP group than in the iliac autograft group. HAP-TCPfillCTfiises directly with the surrounding bonefix)man early stage. TCP also forms a strong union with the surrounding bone by gradual progressive subsUtuUon, and HAP gradually increases in strength after grafting because of its excellentbone conductivity [3]. Thus, the final strength of HAP-TCP filler approaches that of normal cancellous bone. When autogenous bone is used, the graft is gradually substituted by new bone and its mechanical strength remainsreducedduring this process[4]. In the present stucfy, tha^ were no significant differences in age, complications, and non-union period between the HAP-TCP and iliac autograft groups, and there were also no significant difli^Tences in the bone union time or ROM recovery time. CONCLUSIO N There waie no significant differences in bone union time and ROM recov^y time between the HAP-TCP and iliac autograft groups. It was concluded that HAP-TCP was an effective filler for bone defects associated with humeral pseudoarthrosis.

REFERENCES 1. Suzuki, K. and Kurabayashi, H. In: BioceramicsVolume 7, Butterworth-Heinemann, Oxford 1994, 435-440. 2. Suzuki, K., Yamada, M., Yamamoto, K. and Muramatsu, K. In: BioceramicsVolume 8, Pergamon, Oxford 1995, 225-229. 3. Hon, M., Munemiya, M., Takahashi, S., Sawai, K., Niwa, S., Tagai, H., Kobayashi, M., Ono, M. andTakeuchi, K. Cent.Jpn, J. Orthop.Traumat.1984, 27, 2133-2135. 4. Nakamura,S. KitazatoIgaku 1998, 18, 406-419.

Bioceramies, Volume 10 Edited by L. Sedel and C. Rey (Proeeedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd

EXPERIMENTA

L STUD Y OF APATIT E CEMEN T INCLUDIN G CISPLATI N

Y. Tahara, Y. Ishii, S. Sasaki, I. Takano, and K. Ohzeki Department of Orthopaedic Surgery, Kyorin University School of Medicine. 6-20-2 Shinkawa, Mitaka-shi, T(*yo 181, Japan

ABSTRAC T An implant consisting of calcium phosphate cement and cisplatin (CDDP) in order to apply a concept of drug delivery system to the therapy for malignant bone tumor and to maintain an anticancer drug at higher local concentrations with simultaneous complementation of structural disadvantages. In investigation the slow-realing activities in vitro and the influences to other organs and tissues in an implant group and a CDDP systemic dose group in vivo(with Japanese white male rabbits), we concluded that the implant containing 10% CDDP was ideal. KE Y WORD S : Cis-Diamminedichloroplatinum, Ceramic, I>rug Delivery System OBJECTIVE S There are not a few adverse reactions or invasion into oth^ organs and tissues after the existing therapeutic methods for malignant bone tumor. Therefore, we prepared a ceramic implant containing an anticancer drug in order to maintain the anticancer drug at a higher local concentration and simultaneously supplement local structural disadvantages after ^plication of a concept of drug delivery system. Slow-releasing actvities and influences on other organs and tissues were investigated in vitro and in vivo using the implant thus prepared. METHOD S An implant, ISOmg in weight, Smm in diameter and 4mm in height, was prepared with calcium phosphate cement, cicplatin powder and a consolidating solution. Cisplatin was contained in terms weight ratios of 0,5,10 and 20%. For the systemic administration, Randa Injectable manufactured by NIPPON KAYAKU Co., Ltd. containing CDDP by 0.5mg/ml. Japanese white male rabbits weighing about 3kg were used as experimental animals. In vitro slow-releasing experiments, platinum(Pt) in CDDP was determined under the conditions to allow it stand in a thermostat at 37A6 in 100ml of phosphate buffer at pH7.4 at each concentration. In in vivo experiments, each implant was embedded in the distal epiphysis of rabbit femur and changes in body weights and Pt concentrations in the bone marrow surrounding the implant, the bone marrow 1cm distant from the implant, the kidney and the liver were determined. Changes in body weights and Pt concentrations in the bone marrow, the kidney and the liver were determined with a single standard dose of 3mg/kg in humans in a systemic dose group. 345

346


RESULT S In vitro slow-releasin g experiments , with CDD P content s of less than 10%, cumulative slowreleasing rate after 4weeks was about 30% while slow-releasin g velocity was 0.03mg/da y while with a CDD P content of 20% it was 0.1mg/day(Fig . 1 and 2). In vivo the body weight remarkably decrease d in the systemic dose group. With a CDD P content of 20% in which the slow-releasin g velocity and the cumulative slow-releasin g rate were both favorable in vitro, Pt concentration s in the local bone marrow were much higher in the bone marrow and high in other visceral organs and tissues. With a CDD P content of 10%, Pt concentration s in the local bone marrow were higher than the systemic dose group while Pt concentration s in other visceral organs and tissues were lower(Fig. 3,4 and 5).

70 n 60

2 on

I

4)

1

i

50 H

40 30 H

1.0

2.0

3.0

Time(week)

Figurel.Cumulative slow-releasin g rate from the implant.

Figure2. Slow-releasin g velocity from the implant

9000-

—0—

10%Inplant

^

i

— • — 20%Inplant

80007000-

e

1

60005000-

7

4000-

B e u

3000-

a.

200010002

Time(week)

Figure 3. Pt concentratio n in the tissue in a CDD P systemic dose group.

3

4

Time(week)

Figure 4. Pt concentratio n in the bone marrow in an implant -embedding group.

ExperimentalStudy of Apatite CementIncluding Cisplatin: Y. Tahara et al.

347

(fig/tissue-g ) 10-

3

4

Time(week)

Figure 5. Pt concentration in the tissue in an implant-embedding group. DISCUSSIO N The body weight decreased in a wider range in the systemic dose group than in the implantenbedding group so that it might give influences on the gastrointestinal tract. In comparison of Pt concentrations between at the local bone marrow and in the bone marrow 1cm distant, CDDP was considered to be released in a range of less than 1cm. Pt concentrations in the local bone marrow with an implant containing CDDP by more than 5% were more than 100 times higher than those in other visceral organs and tissues so that higher local antitumor effects could be expected under the conditions of few effects on other visceral organs. It is an ideal implant if Pt concentrations are higher locally but lower in other visceral organs and tissues. From the present experiments, an implant containing 10% CDDP is considered as ideal. REFERENCES 1. Uchida A., Shinto Y., Araki N., Ono K., Jpn. J. Cancer Chemother,. 1989,16,3231-3235. 2. Shinto Y., Uchida A., Araki N., Ono K., Jpn. J. Cancer Chemother,.1991,18,221-226. 3. Kitamoto K., Hamanishi C, Yoshii T., Tanaka S., J. Jpn. Orthop. Assoc,. 1994,68,S 1602.


Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) (D1997 Elsevier Science Ltd

IN VIVOEVALUATIO N OF SOL-GE L BIOGLASS* . BIOMECHANICA

L FINDINGS .

Wheeler DL, Hoellrich RG, McLoughlin SW, Chamberland DL, Stokes KE Oregon Health Sciences University, Orthopaedic Research Laboratory 3181 SW Sam Jackson Park Road, L477, Portland, OR 97201

ABSTRAC T Bioglass* (BG) has proven to be an effective bone graft material due to the apatite layer which forms on the surface of the glass, promoting bone formation. Sol-gel Bioglass*, which has greater porosity and surface area, accelerates apatite layer formation and degradability. The objective was to biomechanically evaluate bone formed within distal femoral cancellous bone defects filled with Bioglass*particulates (BG) and two compositions of sol-gel Bioglass* (SGI and SG2) compared to normal cancellous bone (NORM) using a rabbit model. Compressive modulus for the BG group was significantly greater than SGI at 4 and 12 weeks (p cancellous > marrow site. But biodegradation activity was marrow > cancellous > cortical site. After comparison of the two methods, we consider that DXA can bring complementary inquiries in the evaluation of porous ceramics. Its non invasive and atraumatic character should permit in vivo longitudinal survey, and analysis of biodegradation in resorbable ceramics and bone rehabitation in non resorbable ceramics. KEYWORDS : Bioceramics, Bone, Dual energy absorptiometry (DXA), Histomorphometry. INTRODUCTIO N During last years, biomaterials based on calcium phosphate ceramics have been extensively used as osseous substitute in experimental trials and human clinic. This development for clinical practice have necessitated non invasive evaluation. There is numerous methods for postoperative biomaterials measurement. Histomorphometry is helpfiil but traumatic. Conventional imaging modalities are imprecise for quantitation. The aim of this study was to test if Dual energyX-ray absorptiometry (DXA) could bring quantitative evaluation in bone biomaterials in an e?q3erimental model in rabbit, comparatively to histomorphometry. DXA uses a X-ray tube with two different energy levels, enabling correction for the absorption by soft tissue. DXA allows precise, accurate, and non-invasive direct measurement of bone mass in himian and animals ^^’^\early bone changes after total hip arthroplasty ^^\ With adapted software, and high resolution collimators, DXA make possible evaluation of small samples, histologic sections or bones ^’’’^l MATERIA L AND METHOD S Bioceramics:Hydroxyapatite (HA) and Beta Tricalcium Phosphate (P-TCP) with same porosity (50%, measured by Hg porosimeter), pores size (100-300^m) and porous interconnections (30387

388


lOO^m, measured by morphometry), were used. Ca/P ratio was 1.55+0.03 for p-TCP and 1.6 4 –0.03 for HA (measured by Inductive Coupling Plasma). The cylindrical implants were 3 mm diameter (dia.) and 6 – 0.2 mm length. Animals:20 female New Zealand rabbits, 10 month old (adult), 4.0 – 0.3 Kg body weight. Surgical procedure:We used a method deviated from Pasquier et al}^\ In rigorous asepsis conditions and under general anaesthesia, the implants were inserted in a cavity of 3 mm dia. and 6 mm depth in the middle diaphysis (cortico-medullar site) and the extern condyle (cancellous site) of both femurs. Delays: rabbits were euthanazied by overdose anaesthesia initially (TO, 4 rabbits), 12 weeks after implantation (T12, 8 rabbits) and 24 weeks after implantation (T24, 8 rabbits). The total femurs were removed without soft tissue and fixed in 10% neutral buffered formol during two weeks. 8 samples for each ceramic and for each site were removed at T12 & T 24 delays. DXA : In vitroprocedures’.DXA accuracy and reproductibility (precision) were evaluate by comparison with pure weight (PW). 10 HA and 10 p-TCP bioceramics samples were weighted after desiccation. DXA was performed samples soaked in a tissue equivalent material (75^ ethanol, 4 cm depth). We used the manufacturer supplied ultra-high resolution device: line spacing 0.254 mm and point resolution 0.127 mm (Hologic QDR-IOOOAV, Waltham MA USA). Bone Mineral Content (BMC) in mg of equivalent HA or p-TCP was calculated. Ex vivo procedures:The femurs were measured with same technique in vitro,BMC and Bone Mineral Density (BMD) = BMC per projection surface area, mg/mm^ were obtained for regions of interest (ROI) in implanted areas: cancellous (Cn-S) (3.05 x 5.97 mm^), cortical (Ct-S) and medullar (Ma-S) (3.05 x 1.65 mm^), and also in reference (ref) areas: contralateral condyle and proximal diaphysis: (Fig. 1).

D Reference area m litkplantation area

Medullar area D Cortical area

Figure 1.Interest areas for measure Histomorphometr y (HMM) : After DXA analysis, a undecalcified bone preparation was used for each specimen. Residual pores volume (RPV, %), new bone volume (NBV, %), residual material volume (RMV, %) were measured on two 50 ^mi sagittal implant sections with Van Gieson’s Picro-Fuchsine staining. Ratio of material degradation (RMD = 100% - measured RMV / initial RMV) was calculated. Statistical analysis: Results were expressed by means and standard deviations. Impaired bilateral /-test and paired Wilcoxon-test were used to compare the two ceramics, and the bone sites with reference bone by

Reliability of DXA Absorptiometryin Evaluation of Phospho-Calcic Bioceramics:J.X. Lu et al.

389

delay. Correlations between NBV, RMD and BMD were studied with Pearson correlation coeflScient (r). RESULT S AND DISCUSSIO N DXAin vitro(Table1)\ The coefiBcient of variation (CV) of the measure is excellent. Despite narrow range of PW values (from 43.4 to 50.0 mg for p-TCP, 57.0 to 61.5 mg for HA), BMC and PW correlated very well (p.’,ht)>)>yk’//.’,’y

X shrinkage factor prediction sintered specimen

deviation Figure 3. All-ceramic bridge with dense sintered zirconia framework and porcelain veneer. For determining bend strength of bilayered structures as function of the veneer materials’ strength we used zirconia bars of 20 nmi length, 4 mm width and 2 mm thickness. Different veneer materials with MORs ranging from 64 to 270 MPa and elastic moduli of 9 and 70 GPa were applied. Four experimental porcelains (A to D) of similar elastic modulus of 70 GPa and TEC adjusted to zirconia were used. In addition we used one filled polymer (E) with a low elastic modulus of 9 GPa. The total thickness of the specimens was 4 mm. All specimens were tested with the veneer face in tensile mode (first crack).

Figure 2. Test grid specimen to determine shrinkage homogeneity.

RESULT S Process and accuracy In order to evaluate the shrinkage homogeneity we processed dental bridges and test grid specimens as shown in Fig. 2. The linear shrinkage was 22.80 %. The edges of the elevations of the test grid specimen were found not to bulge and remained straight. The average deviation between the predicted dimensions and the sintered array pattern is 19 |im – 12 jim standard deviation for 96 measuring points on the test grid specimen. The dimensional accuracy during 250

,2 CM

O

CI

u

*^porcelain " ^zirconia *-’-"

*^

J

Figure 4. Adjustment of the TEC a of the porcelain to the zirconia (azirconia = 110 [10"^ K*^])-

436


240 T CM

^

200 1

B

D

160 Zr02 veneer veneer porcelain (E = 70GPa) veneer = polymer (E = 9GPa)

120 80

50

100

150 200 250 MO R of veneers [MPa]

300

Figure 5. Bend strength of bilayered structures as a function of the MOR of the veneers. sintering is better than 0.1 % of the total specimen length. This is more than sufficient for dental bridge application. Fig. 3 shows prototypes of an all-ceramic dental bridge with a dense sintered zirconia framework. The dental bridge is placed on the die which demonstrates the high accuracy achieved even with complex free form surface shapes (see Fig. 3 bottom). Veneer porcelain Bilayered structures with a large difference in TEC exhibit a large difference in bend strength for the analysed veneer porcelain thicknesses (see Fig. 4). Specimens with thin porcelain veneers always exhibit higher strength than thick ones. Bilayers with veneer porcelain of similar or slightly lower TEC than zircionia showed comparable bend strengths. For veneering a framework we recommend a porcelain with a slightly smaller TEC than TZP designated by the shaded area in Fig. 4. Bend strength of bilayered structures as function of the MOR of different veneer materials (A, B, C, and D being porcelains with the same TEC of 11.0 x 10"^ K^) and a polymer (E) are shown in Fig. 5. Higher MORs of veneer materials lead to high strength of the bilayers. For lower elastic moduli (i.e. from 70 GPa for B to 9 GPa for E) and similar MOR of both veneers an increasing bend strength of the bilayers is observed. SUMMAR Y A new process "Direct Machining of Ceramics" (DCM) was developed for all-ceramic dental bridges. It comprises ceramic preforms that are machined in a presintered state and subsequently sintered to full density. DCM is feasible for all ceramics. Examples for zirconia (TZP) preforms and complex shaped dental bridges were demonstrated in detail. Final machining in dense sintered state is obsolete due to high accuracy. A veneering porcelain was developed matching zirconia in the TEC behaviour. High strength can be expected for zirconia composites using high strength veneering glass or polymer veneering with low elastic modulus. Further steps comprise the characterisation of all-ceramic dental bridges in laboratory and in vivo. ACKNOWLEDGEMEN T This research is supported by the Swiss Priority Program SPP for Materials Research. REFERENCE S 1. Anusavice, K.J., Ceramic Transactions,1995, 48, 101-124. 2. Liithy, H., in CAD/CIM in AestheticDentistry.,W.H. Mormann, Editor, 1996, Quintessence, Chicago, 229-239. 3. Graule, T.J., Baader, F.H., and Gauckler, L.J., cfi/Ber.DKG, 1994, 71, 4, 317-323.


GRINDIN G OF ZIRCONIA-TZ P IN DENTISTR Y -CAD/CAM-TECHNOLOG

Y FOR

THE MANUFACTURIN G OF FIXE D DENTURES It Luthardt^ W, Rieger^, R, Musil^ ^Friedrich-Schiller-Universitat, Jena, Department of Prosthetic Dentistry, Bachstrasse 18,07740 Jena (Germany), ^Metoxit AG, 8240 Thayngen (Switzerland) ABSTRAC T Due to the rising attention attracted by the biocompatibihty of the material used titan and allceramics-systems are gaining importance. The all-ceramics-systems being currently available are made of glass- and Al203-ceramics. The processing of Zirconia, based on CAD/CAMtechniques, is made possible by the Precident DCS-Systemfi (Ginbach Dental GmbH, Pforzheim, Germany). Regarding its biocompatibihty and material-testing features (flexural strength of 900 MPa compared to 450 MPa of AI2O3) yttrium-oxide-partially-stabilized Zirconia (PSZ) is of great interest as a material for the manufacturing of crowns and bridges [1][2][3][4]. The possibilities of grinding of ceramics such as tetragonal Zirconia-Polycristals (Zirconia-TZP) (Metoxit AG, Thayngen, Switzerland) by the dental technician shall be examined by the presented investigation. KEYWORD S dentistry, CAD/CAM, grinding of Zirconia-TZP, all-ceramics-systems, crowns and bridges INTRODUCTIO N In dental prosthetics distinct alloys are used without any or in combination with ceramics or composites for the manufacturing of fixed prosthesis and implant superstructures. All-ceramics fixed dentures gained interest and importance especially because of their advantages concerning the biocompatibihty of the materials used. Here especially the absence of base alloy-constituents which are necessary to build a bond oxide layer in ceramics fused to metal fixed dentures are of consequence. On the one hand the long-term success of the all-ceramics crowns and bridges is determined by the mechanical features of the frameworks. The resistance against the growing of cracks is especially indicated by the fracture-toughness of the material. So an increased Weibull’s Modulus is more important for the long-term success of all-ceramics crowns and bridges than a high bending strength. On the other hand good mechanical properties are necessary to extend the application of all-ceramics fixed dentures to the posterior teeth or being regarded from a different point of view to increase the safety of the restoration. 437

438


Figure 1: The Zirkonia-TZP-blanks

Figure 2: Sample die of tooth 16

It is not possible to manufacture fixed dentures out of Zirconia-TZP by direct sintering on qualified dies because of the material properties of Zirconia-TZP. Therefore the manufacturing of Zirconia-TZP has to be made by grinding. Grinding- and milling- technology in dentistry require the use of CAD/CAM-Systems in almost all cases. MATERIAL S AND METHOD S The CAD/CAM-System used was the Precident DCS-Systemfi[5]. The material used for this investigation was yttriimi-oxide-partially-stabilized Zirconia [6] manufactured by Metoxit AG [Thayngen, Switzerland]. Figure 1 shows the Zirconia-blanks. Figure2 pictures one of the 6 sample dies of tooth 16 which was used in this investigation. The first step is the definition and recording of the preparation line, the margin of the preparation of the tooth. The digitising of the surface-data of the prepared tooth takes place inside this area. The recording of additional data is possible if the CAD/CAM-system used is able to process this information. Additional data are the contour or the contact surface of the tooth in front of and behind the prepared one. This data may be recorded by optical [7] [8] [9] or mechanical [10] systems. The data-set of the three-dimensional tooth surface will be processed alter the entire recording. The manufacturing-files are processed on the basis of the three-dimensional surface-data together with design parameters, as clearance and thickness of the material. The processed machining files are used to control the CNC-machine. In case of the Precident DCS-Systemfi frameworks out of titan are milled or Zirconia-TZP and InCeram are ground. After the CAD/CAM-manufacturing a dental technician has to adapt theframeworkto the model die [11]. Figure 3 presents one framework of Zirconia. For the investigation of the frameworks the machined blanks were separated, the parting planes were polished and SEM pictures were taken. Later, to avoid side eflFects by sample preparation modified blanks were used. For this Metoxit AG [Thayngen, Switzerland],

Grinding of Zirconia - TZP in Dentistry: R. Luthardt et al.

439

>.<w|i4-_

Figure 3: Framework 16 of Zirconia

Figure 4: Modified blank after grinding the inner surface of a framework

produced half-cylinders . The margins were bevelled , the parting planes polished and two halfcylinders adhesivel y bonded. This new cylinder was poured in epoxy-resin . Both the blanks and the modified blanks were machined with various grinding parameters using the Precident DCS-Systemfi. Figure 4 shows one half of a modified blank after grinding the inner surface of aframeworkand before SEM-examination. TH E SEM-EXAMINATIO N Microcracks reduce the long-term success of all-ceramics crowns and bridges. The more and the longer the cracks the worse are the mechanica l properties . So CAD/CAM-manufacturing of ceramics has to avoid the developmen t of cracks [12]. The SEM-examination should proof the presence of grinding-relate d rim-damage and microcracks.

Figure 5: Locations where the SEM-pictures were taken

440


The points within the blanks where the SEM-pictures were taken are depicte d by figure 5. These are the inner top surface, the side of theframeworkand the area of the preparation-line . Under the aspect of manufacturing engineerin g the worst grinding relations are to be found on top of the inner surface. The side allows answers to the surface condition to be reached. The area of the preparation-lin e has to stand the highest force while chewing but has the smallest thickness of the framework. CONCLUSIO N The results of the examinatio n show the possibility to manufacture fitting fixeddentures using the Precident DCS-Systemfi, These results are confirmed by the SEM-pictures of the ground surface. To use CAD/CAM-Technology instead of the casting of alloys it is necessar y to shorten the manufacturing time of the ceramics. If the manufacturing time of fixed dentures can be decreased , the expansion of the application of all-ceramics-system s will be ob\ious.

REFERENCES: [I]

[2] [3] [4] [5] [6] [7] [8] [9] [10] [II ] [12]

Christel, P. et al: Mechanical propertie s and short-term in-vivo evalation of yttriumoxide partially-stabilize d Zirconia. J Biomed Mater Res 1989, 23, 45-61. HolscKW.,KapperU H. F.: Festigkeitsprufim g von Vollkeramische n Einzelzahnersat z fur den Front- und Seitenzahnbereic h Dtsch zahn^rztl Z 1992, 47, 621-623 . RiegerW.:Aluminium- und Zirkonoxidkeramik in der Medizin ID R 1993, 2/93, 116-120 . Stevens, K: Zirconia and Zirconia Ceramics, An introductio n to Zirconia Magnesium Elektron Publikation No. 113, Second edition Litho 2000, Twickenham/UK 1986. Luthardt, R Musil,R,: Das Precident DCS-Systemfi -CAD/CAM-gefertigter Zahnersatz aus Titan und Zirkonoxid Phillip 11996,13,217-225 . Maier,K R,: Leitfaden technisch e Keramik; WeikstoflBkunde II , Keramik, Selbstverlag Institut fiir keramische Komponenten im Maschinenbau, Aachen 199L Benz,C , Schwarz, P.: Wie genau ist der optische Cerec-Abdruck? Dtsch zahnarztl Z 1991, 46, 632-634 . Bose,M , Ott,K. K R.: Wissenschaftlich e Erkenntnisse tiber das Cerec-Syste m Dtsch zahnarzU Z 1994,49 , 671-673 . Zel,vanderJ, M ; C AD/C AM-Restaurationen in der Okklusion ZahnarzU Welt 1994, 103, 420-425 . Bieniek, K. W.:Computer- und andere automatisiert e Systeme zur Erstellung von Zahnrestauratione n ZM K 1994, 5/94, 6-14. Holmes, J.R.,et al: Considerations in measurement s of marginal fit J Prosthet Dent 1989,62 , 405408 . Marx,R,: Modeme Keramische WeiiestofiF e ftir ^sthetisch e Restauratione nVerstariamg und Bruchzahigkeit Dtsch zahnarzU Z 1993, 48, 229-236 .

Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris. France, October 1997) '1997 Elsevier Science Ltd

ZmCONI A IMPLANT S WIT H A PLASM A SPRAYE D SiOj-HA BIOACTIV E COATIN G Alexandre Pedra, Patrick Sharrock Rimimplant, Reims and Laboratoire de Chimie Inorganique, Universite Toulouse III, France. KEYWORD S Zirconia, dental implants, coating, Hydroxylapatite-silica . ABSTRAC T Yttria stabilized zirconium oxide is used for its high mechanical strength to manufacture the core of a dental implant. Bioreactivity relies on a plasma-sprayed undercoat containing zirconia and silica and an exterior silica-hydroxylapatite surface layer. Histological and clinical evaluations demonstrate excellent osteointegration. Load bearing implants are functional. Good initial clinical results are important for long term serviceability and success rates. INTRODUCTIO N Tetragonal zirconia, ZrOz, stabilized with Y2O3 is attracting much attention as a high quality technical ceramic (1-4). The enhanced mechanical properties of zirconia ceramics characterized by a flexural strength near 900 MPa and a toughness of lOMPa makes it an ideal biomaterial for the manufacture of dental implants (5-6). Crystalline calcium phosphates, hydroxylapatite (HA) or tricalcium phosphate (TCP) are too brittle to be considered for the manufacture of load-bearing implants (7). This has led to the development of HA plasma-sprayed coatings (8). Failure to obtain an adequately adherent HA bioactive coating on the ceramic substrate led us to examine various undercoats on zirconia in order to obtain a functional gradient. The Prosiap (projected silica and apatite) layer was successful and resulted in the manufacture of the first zirconia based dental implant with bioactive coating to be marketed in France in 1993. We report here the physico-chemical characteristics of the RIM implant and some features of the clinical results obtained with prostheses in use on the ceramic implants. MATERIAL S AND METHOD S RIM implants were manufactured by machining PSZ green bocfy cylinders followed by sintering (ZrOz with 3mol % Y2O3 from Ceramiques Techniques Desmarquest) (9). The implants have a cylindro-conical shape with three lengths (13, 16, 19 mm.) and widths (2.8, 3.3, 3.8 mm.). The transgingival collars are polished, while the endosseous root portions are coated with the Prosiap deposit. X-ray patterns were studied directly on coated samples. A few grains were detached from the coating for FTIR spectroscopy. Seven adult rabbits were used for the in vivo experiments. The endosseous parts of the implants were cut off and sterilized by gamma radiation. Two implants were impacted in each tibial epiphysis area in various locations in order to study the biological response in medullar, spongy and cancellous areas. Rabbits were sacrificed at 3, 6, and 12 weeks post-operatively and non decalcified sections obtained by infiltrating, cutting grinding techniques. RIM implants were placed in patients requiring dental root replacement using classical progressive drills cooled with internal serum irrigation, and a final drill with an adapted conical shape and calibrated depth markers. 441

442


3000

Figure 1. X-ray pattern of Prosiap before(a) and following immersion (b).

Figure 2.1.R. spectra of Prosiap before(a) and following immersion(b).

RESULTS AND DISCUSSION The Prosiap coating consists of a Zr02-Si02 plasma-sprayed deposit of 20 um thickness covered by an HA-SiOi layer 30 um thick.The resulting zirconium containing silica melt provides an adherent undercoat for the subsequent HA outer layer. Mechanical tests performed on cylindrical test-pieces yielded lower tensile strength values for HA (12.7 MPa) than for the undercoat (61.8MPa). In the zircon-zirconia system the phases are compatible in all proportions at temperatures up to 1675 C (10), which may explain the strength of the bonding layer. The HA+undercoat (56 Mpa) and silica-HA+undercoat (57 Mpa) both have good tensile strengths which may be explained by glass fiision and chemical bonding by calcium silicate formation at interphases.The X-ray diffraction spectrum of the undercoat shows large and strong reflections at Cu Kalpha (theta) values of 15 and 25 , indicative of tetragonal zirconia in the substrate with typical d values of 2.968 and 1.821 A (Angstroms). Smaller peaks at d = 3.162 and 2.840A reveal the presence of monoclinic zirconia in the bonding layer. The plasma-sprayed silica-HA coating shows an X-ray pattern illustrated in figure la. The coating transforms into a more typical HA following immersion in Ringer’s solution during one week, as shown in figure lb. There is a small but perceptible shift in the position of the main HA peak from d == 2.8136A in the deposit to d = 2.8122A in the recrystallized deposit, accompanied by a doubling of the intensity. The starting HA powder has d = 2.8117A, shifting to d = 2.8141 A after mixing withfilmedsilica and calcining at 900 C. The infrared spectra confirm the amorphous nature of the deposit and its tendency to recrystallize when immersed into Ringer’s solution. Figure 2 illustrates the corresponding changes. The absorption peak at 1650cm’^ in figure 2b is related to carbonates present in the recrystallized coating. The Si-O-Si vibrations are absent from the 2b spectrum suggesting the HA originates in fact from new crystals deposited on the coating. All implants demonstrated good osteointegration at all time periods. By three weeks a periosseous formation was present surrounding the implants and sealing the cortical implant access zones. A thin 50 um wide, layered bone apposition followed the implant surfaces in the medullar areas. Figure 3 illustrates this new bone -bioactive coating interface- No spalling of the coating was observed. At six weeks, neoformed bone was attached to the implant in cortical situation and followed the surface geometry of the implant into spongy areas. Signs of bone remodelling near the implants were visible, with no deposition offibrousencapsulation tissue. At

Grinding of Zirconia - TZP in Dentistry: R. Luthardt et al.

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Figure 3. X-ray view of two RIM implants in rabbit epiphyseal bone at six weeks. twelve weeks the implants are very firmly bound, with endosteal bone proliferating along the ends of the implants protruding in the tibiae. The bone marrow had a normal appearance and we could not detect any significant dissolution or resorption of the coating. The clinical uses of RIM implants have been described and are multiple (11). Partially and totally edentulous patients have been treated successfully. Extraction-immediate implantation has been performed as well as implantation with simultaneous bone augmentation with HA products. These procedures are usually hazardous and difficult with bare metal implants, but seen

Figure 4. Retroalveolar view of 4 RIM implants placed in the maxilla. Note the natural bone height 26 weeks post-implantation and the absence of visiblefibrousinterposition.

444


to be well tolerated with Prosiap coated zirconia implants. The example illustrated here concerns a woman with seven implants placed simultaneously in January 1995. Following implant integration, temporary crowns were fitted which were replaced by permanent, individual, sealed ceramic crowns on standard inlay cores. An adjacent revitalized adjacent tooth was also treated. None of the implants were connected. The patient was checked regularly. The panoramic view in November 1995 confirmed the absence of periapical infection or peri-implant inflammation. In March 1997 retroalveolar X-rays revealed good bone behaviour 26 months post-implantation, without any interposition of fibrous tissue. Ceramics are well accepted physiologically and psychologically as well. Overall patient satisfaction is excellent CONCLUSIO N Stabilized zirconia ceramics provide adequate mechanical strength for load bearing dental implants. HA-silica bioactive plasma-sprayed deposits with a proper ceramic bonding imdercoat stimulate early bony integration. The one step gingival procedure allows simultaneous soft tissue healing and yields very good esthetic results and tight gingival seals effectively preventing peri-implant inflanunation and encouraging patient compliance with proper hygiene. We feel technical ceramics combined with a bioactive coating provide an ideal material for challenging dental reconstructions.

RERERENCES 1. H. Liu, Q. Xue, investigation of the crystallization of Zr02 (Y2O3 3 mole%) nanopowders, J. Mat. Res (1996), 917-921. 2. W. Burger, H.G.Richter, C.Piconi, R.Vatteroni, A.Cittadini, M.Boccalari, New Y-TZP powders for medical grade zirconia, J. Mater. Sci.,Mat. in Med., 8, (1997), 113-118. 3. C.L Curtis, D.T.Gawne, M.Priestnall, The processing and electrical properties of plasmasprayed yttria-zirconia, J. Mat. Sci, 29, (1994), 3102-3106. 4. M.Mattioli-Belmonte, P.Mengucci, N.Specchia, G.Golbi, S.Dubini, L.Simonelli, F.Greco, G.Majni,G.Giagini,C.Rizzoli, An experimental study in X-ray spectroscopy of the zirconium (Ca-PSZ) bone interace. Microanalytical evaluation of the osteogenic response, J. Mat. Sci., Mat. in Med, 8, (1997), 85-90. 5. A.H.Heuer, L.W.Hobbs, Science and technology of Zirconia, Advances in ceramics vol 3 American Ceramic Society, Westerville, Ohio,(1981). 6. M.S.Zolotar, C.A.C.Zavaglia, Fracture toughness and microstructure degradation of Y-TZP in aqueous physiological environnement, J. Mat. Sci. Mat in Med., 7, (1996),367-369. 7. J.Li, L.Hermansson, R.Soremark, High strength biofimctional Zirconia: mechanical properties and static fatigue behaviour of zirconia apatite composites. , J. Mat. Sci. Mat in Med., 4, (1993),50-54. 8. Y.Kawamoto, Y.Yokogawa, M.Toriyama ,S.Kawam\u^, T.Suzuki, Coating of Beta-tricalcium phosphate on Yttria-partially stabilized zirconia using magnesium metaphosphate as an interlayer. J. Ceramic. Soc. Jap. Int-ed, 99, (1991), 19-22 9. B.Cales, High reliable zirconia ceramics for orthopaedics. Fifth world biomaterials congress, Toronto, 1996, 177. 10. T.Itoh, Zircon ceramics prepared from hydrous zirconia and amorphous silica. J. Mat. Sci. Let, 13, (1991), 1661-1663. ll.J.Pedra, P.Sharrock, Un nouvel implant dentaire bioactif en zircone: I’implant RIM, Actualites en Biomateriaux, 4, (1995), 405-410 .

ORTHOPAEDIC APPLICATIONS


Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedings of the10thInternational Symposium on Ceramics in Medicine, Paris,France,October1997) '1997 Elsevier Science Ltd

EFFEC T OF TIM E AND TEMPERATUR E ON THE PRODUCTIO N OF POROU S ELECTROLYTI C HYDROXYAPATIT E COATING S N. Asaoka*+, S. Best+ and W. Bonfield+ *Central Research Institute, Mitsubishi Materials Corporation, 1-297 Kitabukuro-cho, Omiya, Saitama, 330, Japan +IRC in Biomedical Materials, Queen Mary and Westfield College, Mile End Road, London El 4NS, U.K.

ABSTRAC T Porous hydroxyapatite coatings were produced on pure titanium substrates by the electrolysis of clear solutions containing calcium and phosphate ions and complexing agents at a constant DC voltage. The effect of the electrolytic conditions, i.e. the temperature, duration and applied voltage, on the properties of the coatings were investigated. The coatings were analysed using Xray diffraction (XRD) and scanning electron microscopy (SEM). No deposit was obtained at applied voltages higher than 5V. The coatings prepared at temperatures higher than 60T were identified, using XRD, as low crystallinity hydroxyapatite and SEM revealed that they had a porous honeycomb-like structure. The pore size varied from 0.5 to 2 microns depending on the temperature at which electrolysis was performed. KEYWORD S hydroxyapatite, titanium, coating, electrolysis, complexing agent INTRODUCTIO N Hydroxyapatite is one of the most widely researched bioactive ceramics. However, its clinical use has been restricted to non-major load bearing applications due to its relatively inferior mechanical properties. In order to take advantage of the bioactivity of the material and to overcome the disadvantages of mechanical performance, a number of methods have been investigated of applying hydroxyapatite coatings to metallic substrates. One of the most popular techniques currently in use is plasma spraying [1] and the vast majority of commercially manufactured hydroxyapatite coated implants for dental and orthopaedic applications are produced using this route. However, there are some disadvantages with the use of thermal spray techniques such as phase decomposition, poor coating reproducibility and debonding at the coating/substrate interface. A variety of alternative techniques, are therefore under investigation including dipping and sintering [2], sputtering [3], HIP [4], electrochemical deposition [5] and electrophoresis [6], although, currently there do not appear to be any commercially viable techniques torivalplasma spraying in terms of mechanical strength or productivity. A process for the production of uniform monetite coatings on titanium substrates by electrophoresis in clear solutions has been reported previously [7]. In this study, results are reported from a new process to prepare porous hydroxyapatite coatings directly from solution. 447

448


Figure 1 Experimental apparatus for electrolysis MATERIAL S AND METHOD S Reagent grade orthophosphoric acid, calcium hydroxide and a complexing agent were dissolved in distilled water to prepare clear electrolytic solutions. Ammonium hydroxide solutions were utilised to adjust pH. The electrode set was composed of a commercial purity titanium plate (50x20x1 nmi^) as a cathode and a pure platinum plate (50x20x0.125mm^) as an anode. The titanium and platinum substrates were degreased using acetone and were then placed parallel with each other at a separation of 10mm. The electrodes were dipped into a vessel containing \QOcvc? of the clear electrolyte at a depth of 15mm for electrolysis (Fig.l). When the solution reached a predetermined temperature (from 20 to 80T) a constant DC voltage up to 10 volts was applied between the electrodes. As electrolysis commenced, the alkaline solution was dripped into the vessel and the electrolyte was slowly stirred to be homogenised. The current and pH of the electrolyte were monitored during electrolysis. After electrolysis for up to 60 minutes, the electrode sets were gently rinsed in distilled water to remove excess electrolyte and dried at room temperature for several days. Then the dried coated substrates were weighed to calculate the net weight of the coating layers. Phase identification was made using X-ray diffraction (Siemens D50(X)) with a tube voltage 40kV, 40mA and scanning rate 3.2deg(2theta)/min.. The structure of the coatings were observed using a scanning electron microscope (JEOL JSM-6300F) with accelerating voltage 5keV. RESULT S AND DISCUSSIO N The initial pH of the orthophosphoric acid, calcium hydroxide and a complexing agent mixture was less than 4 and on addition of the ammonium hydroxide solution, the pH increased up to a final value of pH 8 with precipitation commencing at above pH 5. There was a marked effect of voltage on coating formation: at 60"C, the net weight of the cathode increased with increasing voltage up to 5V, but at voltages above this, no deposition occurred (Fig.2). Generally, the coating weight increased with increasing temperature (except for the reaction performed at SOT at 5V). The fact that no coatings were obtained on the cathode with voltages higher than 5 volts is possibly due to the vigorous bubbling resulting from the electrolysis of water molecules which was observed on the surface of the both electrodes. In this case there appeared to be more

Effects of Time and Temperatureon Production of Porous ElectrolyticHA Coatings: N. Asaoka et al.

449

electric power consumed in the electrolysi s than in depositio n and as a result, coating formation may have been prevented . The bubbling was associate d with a correspondin g increase in electric current. As shown in Figure 3, the current increase d rapidly with increasing applied voltage and temperature . This may have affecte d the physical adhesion of the coating layers, and the sudden decrease of the coating weight at 80"C and 5 volts in Figure 2 may also be explaine d by the partial decompositio n of the layer by the bubbling. Electrolysis temperatur e also affecte d the formation and appearance of the coatings: the colour of the coatings prepared at 60X or above, was pure white while those prepared at lower temperature s were light grey. However, extendin g the duration of electrolysi s did not necessaril y result in an increase in coating weight. Figure 4 indicates that coating weight increase d from 10 to 20 minutes, while it decrease d at 30 minutes. The results from X-ray diffraction indicated that the coating layers deposite d at a temperatur e of 60*C or higher were composed of low crystallinity hydroxyapatit e while those deposite d at temperature s lower than 60"C were mainly composed of brushite (CaHP04-2H20). 10 E S "S

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Figure 4 Effect of electrolysi s duration on the coating weight at 60T, 5V

Figure 5 Porous hydroxyapatit e coating surface (SEM image), prepared at 60**C and 5V for 10 minutes. )

450


At the beginning of precipitation, the pH of the solutions was about 5. Therefore the phase composition of the initial precipitates and the coatings would have been brushite which is more stable than hydroxyapatite at lower pHs. As the pH continued to increase throughout the reaction, hydroxyapatite would have become more stable. However, an activation energy would have been needed to be supplied to convert crystallised brushite to hydroxyapatite and it is likely that, only at relatively high temperatures, would this activation energy have been sufficient. SEM observation revealed that the hydroxyapatite coating layers were uniform with a porous honeycomb-like structure. The pore sizes appeared to be more strongly influenced by the reaction temperature than reaction duration and, as indicated in Figure 5, the average pore size for a reaction performed at SOT (3V) was approximately 2 microns where for a coating deposited at 6 0 T the average pore size was approximately 0.5 microns. The porous hydroxyapatite coatings observed in this study are in contrast to monetite coatings, prepared previously [7] which were reported to consist of densely packed uniform plate-like particles. This structural difference may reflect the morphological preference in crystal growth under the applied electric field. The hydroxyapatite coatings prepared at SOT had a larger pore size than those at 60X. This difference seemed to be dependent on the growth rate of the crystals with precipitation temperature. CONCLUSIO N Hydroxyapatite coatings can be prepared using an electrolytic process in solution containing calcium and phosphate ions with a complexing agent. At applied voltages of 5 volts or lower a fine white deposit resulted on the cathodic titanium plate. Coatings prepared at 6 0 T or higher consisted of low crystallinity hydroxyapatite, while brushite coatings were obtained at lower temperatures. The hydroxyapatite coatings had a very porous structure with pores of 0.5 to 2 microns in diameter. ACKNOWLEDGMENT S The support of the EPSRC for the IRC in Biomedical Materials and of Mitsubishi Materials Corporation for one of the authors (N.A.) are gratefully acknowledged. REFERENCE S 1. de Groot, K., Geesink, R., Klein, C.P.A.T.and Serekian, P., J. Biomed. Mater. Res., 1987, 21, 1375-1381 2. Yankee, S.J., Pletka, B. J., Luckey, H. A. and Johnson, W. A., Therm. Spray Res. Appl., 1991,433-438 3. Ong, J. L., Lucas, L. C , Lacefield, W. R. and Rigney,E. D., Biomaterials, 1992, 13 ,249254 4. Heroe, H., Wie, H., Joergensen, R. B. and Ruyter, I. E., J. Biomed. Mater, Res.,1994, 28, 343-348 5. Shirkhanzadeh, M., J. Mater. Sci. Lett., 1991,10, 1415-1417 6. Ducheyne, P., Radin., S., Heughebaert, M and Heughebaert, J. C , Biomaterials, 1990,11, 244-254 7. Asaoka, N. Best, S. M. and Bonfield, W. In Bioceramics Volume 9, Elsevier Science Ltd., Oxford, 1996, 289-292


CALCIU M PHOSPHAT E FORMATIO N ON CHEMICALL Y TREATE D VACUU M PLASM A SPRAYE D TITANIU M COATING S S.-W. Ha^ K.-L. Eckert^ H. Gnrner^, E. Wintermantel^ ^ Chair of Biocompatible Materials Science and Engineering, Department of Materials, ETH Zurich, Wagistr. 23, CH-8952 Schlieren ^ Medicoat AG, Gewerbe Nord, CH-5506 Magenwil ABSTRAC T Carbon fiber reinforced polyetheretherketone (CF-PEEK) substrates were coated with titanium (Ti) by vacuum plasma spraying (VPS) and chemically treated in lOM sodium hydroxide (NaOH) or 30% hydrogen peroxide (H2O2) solution. After chemical treatment, the specimens were immersed in simulated body fluid (SBF) containing ions in concentrations similar to those of human blood plasma. Scanning electron microscopy (SEM), energy dispersive X-ray analysis (EDX) and diffuse reflectance fourier transformed infi*ared spectroscopy (DRIFT) were used to analyse the chemically treated VPS-Ti surface and the calcium phosphate layer formed during immersion in SBF. It was observed that a carbonate containing calcium phosphate layer was formed on the NaOH treated VPS-Ti surface during immersion in SBF, whereas no calcium phosphate precipitation occurred on the untreated and H2O2 treated surfaces. It is therefore concluded that vacuum plasma spraying with Ti and subsequent chemical modification in 10 M NaOH solution at 60 C for 2 hours is a suitable method for the preparation of bioactive coatings for bone ongrowth on CF-PEEK. KEYWORD S Vacuum plasma spraying, carbon fiber reinforced PEEK, titanium, calcium phosphate coatings INTRODUCTIO N Vacuum plasma spraying on CF-PEEK is currently being established to produce Ti and hydroxyapatite coatings in order to make this thermoplastic composite suitable as implant material [1]. It is assumed that such coatings will enable long-term fixation of this material in bone tissue. In the present study, two different methods, NaOH treatment or H2O2 treatment on VPS-Ti surfaces were peformed to provide nucleation sites for calcium phosphate formation in a simulated body fluid. The characteristics of the modified surface and the induction of in vitrobiological apatite for› mation were investigated. MATERIAL S AND METHOD S CF-PEEK (Ensinger GmbH, Germany) disks with a diameter of 10 mm and a height of 7 mm were sandblasted with alumina, cleaned with ethanol and deionized water and dried in a vacuum oven at 200 C for at least 7 days. VPS was performed with fine Ti powder (d5o = 25 |Lim) and coarse Ti powder (d5o = 120 |im). Chemical treatment of the VPS coatings was performed in 30% H2O2 at room temperature and in 10 M NaOH solution at 60 C. Soaking time was 2 hours for both treat› ments. After chemical treatment the specimens were washed in distilled water and immersed in sim451

452


ulated body fluid (SBF) which was prepared as described in [2]. Immersion was carried out in 25 ml of SBF during 1, 4, 10 and 24 days in a laboratory shaker (Infors AG) rotating with 80 rpm at 1>TC and pH 7.4. After immersion the specimens were gently rinsed in distilled water and dried in air. Scanning electron microscopy (SEM) and energy dispersive X-ray (EDX) analysis was performed to characterize morphology and chemical composition of the coatings. The specimens were coated with platinum in a sputter coater before SEM and EDX analysis. Chemical changes of the substrates after immersion in SBF were examined by diffuse reflectance fourier transformed infrared (DRIFT) analysis (Perkin Elmer System 2000). Infrared spectra were obtamed in a wavenumber range of 4000-400 cm 1 RESULT S AN D DISCUSSIO N In fig. 1, the SEM micrographs of untreated, H2O2 treated and NaOH treated VPS-Ti surfaces are shown. After NaOH treatment the formation of a fibrous, needle-like structure was observed. No morphological changes were detected on the untreated and H2O2 treated surfaces. EDX spectra showed that Na was incorporated into the VPS-Ti coating after NaOH treatment (fig. 3 left) and it is assumed that the newly formed structures as observed with SEM contain sodium. Besides Na and Ti no additional elements were found on the NaOH treated samples. No changes in chemical composi› tion were observed on both, untreated and H2O2 treated surfaces. After immersion in SBF, the as-received and the chemically treated VPS coatings were analysed with EDX and SEM. Fig. 3 shows the EDX spectra of the untreated and treated VPS-Ti coatings. In both, untreated and H2O2 treated substrates, only the Ti peak coming from the VPS-Ti coating and a little amount of platinum (Pt) from the sputter coating was detected. SEM analysis revealed that no calcium phosphate layer was formed on the as-received and H2O2 treated VPS-Ti surfaces after 24 days of immersion in SBF (fig. 2). On the NaOH treated coating, formation of calcium phosphate was observed aheady after one day of immersion in SBF. After 24 days the VPS-Ti surface was completely covered with calcium phosphate precipitates (fig. 2 and 3). Fig. 4 shows the FTIR spec› tra of the NaOH treated VPS-Ti surfaces, which were immersed in SBF for 1, 4 and 24 days, show› ing the continous growth of a carbonate containing calcium phosphate layer, which is regarded to be similar to the chemical composition of biological apatite m the natural bone [3-5].

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Calcium PhosphateFormationon ChemicallyTreatedVPS TitaniumCoatings: S.- W. Ha et al. 453

The results of the present study demonstrated that H2O2 treatment of VPS-Ti surfaces does not induce the formation of calcium phosphate. After NaOH treatment a small Na peak appeared in the EDX spectra. This small amount of Na could result from the formation of sodium titanate in accor› dance to the observations reported in [6]. In that work the formation of Na2Ti50ii on various pol› ished Ti alloy surfaces after alkaline treatment in NaOH was shown. During immersion in SBF calcium phosphate formation could have occurred due to a ion-exchange process, since sodium titanate shows a high ion-exchange capacity and all its Na^ ions are exchangeable [7].

untreated

Figure 2

NaOH treated

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454

Bioceramics Volume10 1

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FTIR spectra of VPS-Ti coatings on carbon fiber reinforced PEEK after NaOH treatment and subsequent immersion in SBF at 37 C, obtained in the diffuse reflectance mode. With increasing immersion time, phosphate and carbonate bands became more distinct indicating the continuous precipitation of a carbonate containing calcium phosphate.

CONCLUSIO N The present study has demonstrated that vacuum plasma spraying of Ti and subsequent chemical modification in 10 M NaOH at 60 C for 2 hours is a suitable method for the preparation of bioactive coatings on carbon fiber reinforced PEEK. Carbonate containmg calcium phosphate, which has nearly the same morphology and chemical composition as biological apatite in natural bone, has been formed on the NaOH treated VPS-Ti coatmgs durmg immersion m SBF. The VPS-Ti coatings showed a very rough topography and a high surface area. This is assumed to positively affect the chemical modification of the Ti surface by NaOH treatment and the in vitrocalcium phosphate dep› osition on the modified VPS-Ti surface.

REFERENCES [1] [2] [3] [4] [5] [6] [7]

Ha S.-W., et al., in: Bioceramics,Vol. 10, Elsevier Science Ltd., Oxford, UK, 1997, in press. Kokubo T., et al, in: Bioceramics,Vol. 4, Butterworth-Heinemann, Guildford, 1991, 113-120. Maruno S., et al, Biomaterials,12, 1991, 225-230. Brophy G.P., and Nash J.T., The americanmineralogist,53, 1968, 445-454. Heughebaert M., et al., Journal of BiomedicalMaterialsResearch,22, 1988, 254-268. Krni H.-M., et al. Journal of BiomedicalMaterialsResearch,32, 1996, 409-417. Clearfield A., et al.. Journal of solid statechemistry,73, 1988, 98-106.


PROPERTIE S OF PLASM A SPRAYE D BIOACTIV E FLUORHYDROXYAPATIT E COATING S RANZ X.l, REY C.l, ANTOLOTTI N.2, HARMAND M.F.3, MORONI A 4, ORIENTI L 4, VIOLA 0.2, BERTINI 5.2. SCRIVANI A 2 1 Laboratoire dc Physico-Chimie des Solides INPT-ENSCT. CNRS-UPRESA 5071, 38 rue des 36 ponts, 31400 Toulouse, France 2 Hametal S.p.A. Biocoatings div.. Via G. di Vittorio 51,43045 Fomovo Taro, Italy 3 LEMI, Technopole Montesqieu, 33650 Martillac, France 4 Istituti Ortopedici Rizzoli, via Di Barbiano 1/10,40136 Bologna, Italy ABSTRAC T In order to improve both the resistance to in vivo biodegradation and the osteoconductivity of plasma sprayed coatings, we developed a bi-layer coating composed of a stable fluorhydroxyapatite plasma sprayed coating covered by a poorly crystalline carbonate sâtite layer analogous to bone mineral. In vitrotests showed that despite a very slight cytotoxicity, the presence of the carbonate layer improves osteoblast proliferation and colonisation. The bi-layer coating bone ingrowth and bone attachment behaviour is similar to the more soluble hydioxyapatite plasma sprayed layers. KEYWORD S Plasma sprayed coating,fluorhydrohyapatite,solubility, surface treament, osteoconductivity. INTRODUCTIO N Plasma sprayed hydroxyapatite (Caio(P04)6(OH)2: HA) coatings on titanium alloys are largely used on orthopaedic prostheses owing to their ability to enhance bone formation and to establish strong bone/implant bonding [1]. Unfortunately, HA decomposes during the plasma spraying process into more soluble compounds [2,3] which can induce release of debris and detachment of the coating [4]. On the other hand, the osteoconductivity of Ca-P coatings has been shown to be related to their partial solubility [5]. It has been reported in a previous paper that fluorhydroxyapatite coatings (Caio(P04)6FOH: FHA) decompose less than HA during plasma spraying and lead to more adherent coatings [6]. In this report we compare the dissolution of HA and FHA plasma sprayed coatings at constant pH, and we describe a surface treatment allowing the formation of a carbonateapatite layer on the FHA coating. The main chemical-physical characterisatics of this layer (FTER, X-ray diffraction, SEM, specific surface area) are given and its biological activity determined by in vitroand in vivo studies. MATERIAL S AND METHOD S All the plasma spray depositions of stoichiometric HA and FHA were performed by Biocoatings. The kinetics of dissolution was measured at constant pH and temperature (pH=4 and T=37X) in lactic acid solution. The dissolution rate of HA and FHA coatings were determined by the quantity of acid required to maintain constant pH. After 20 hours the samples were dried and weighed then, they were sonicated to release loose particles and weighed again; this test gave an indication of the disaggregation of the coating. 455

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Poorly crystalline carbonate apatite was precipitated on the surface of FHA from supersaturated solutions containing Ca(NC)3)2, (NH4)2HP04 and NaHCOs [7]. The in vitrotests WCTC performed according to European and French standards (ISO 10993-3; NF S 91-145; NF S 91-142). The in vivo implantation into sheep and the biomechanical behaviour and histological obsCTvations were carried out according to published procedures [8]. RESULT S AND DISCUSSIO N Substrate characterisatio n FHA has been shown to be one of the least soluble Ca-P apatites [9]. The physical-chemical characterisation of the two coatings is reported in table 1 [6]. Constant pH and temperature dissolution curves are represented in figure 1. For coatings with similar specific surface areas (around 0.55 m^.g-^), the data indicate a slower dissolution rate for FHA than for HA.

Samples HA FHA

Table 1. Characteristics of HA and FHA coatings Phase composition (%) Physical charactoistics Oxy. Amo. P-TCP TTCP Ca/P Density Porosity Crystal. (atomic) (m^g-l) (%) (%) 65 20 1.0 1.1 1,667 3.08 15.3 83 07 16 0.4 1.8 1,666 3.13 9.8 81

Although it is difficult to relate the kinetics of dissolution of the coatings to a precise cause, the presence of foreign phases more soluble than âtite, especially amorphous calcium phosphate, appears as the main parameter (table 1) [2, 3]. Moreover, fluoridation of apatites is known to improve their resistance to acid dissolution [11]. However, dissolution is not the only mechanism involved in the biodegradation of Ca-P coated implants; disaggregation of the coatings, which results in the release of fine particles of apatite into the solution, also occurs. AftCT dissolution and sonication, we measured the loss of weight (Fig. 2). Disaggregation was seen to be lower for FHA than for HA. The parameters are essentially the same as for dissolution. A predominant role of the amorphous Ca-P phase can also be suggested. This phase is among the most soluble of the Ca-P compounds and behaves like a binder between crystalline particles. Furthermore, hydrolysis of CaO into Ca(0H)2 favours crack formation inducing the disaggregation of the Ca-P layer [12]. FHA gives less soluble and more resistant Ca-P plasma spayed coatings. If we consider that bone remodelling in the vicinity of the implant involves acid dissolution, it may be expected that FHA coatings will last longer than HA coatings. On the otherhand, the osteoconduction seems to be related to the solubility of the material [5]. The biological activity of FHA might then be limited although FHA seems to favour cell proliferation and to increase bone density [14]. Surface treatmen t The aim of is this treatment was to form a poorly crystalline apatite at the surface of the coating to improve the biological activity. After treatment, SEM analysis showed the presence of small sph^cal particles from 0.5 to 2 ^m in diameter covering the entire surface of the substrate (Fig. 3). The chemical composition determined by EDS indicated that calcium and phosphate were associated with sodium and carbon (carbonate). Furth^more the specific surface area increased by about 12 % after treatment due to the presence of small crystals on the substrate. The X-ray diffraction patton of the precipitate famed by homogeneous nucleation, in the same conditions, revealed a poorly crystalline apatitic phase (Fig. 4a). The average dimension of crytals det^mined from band broadening of the 002 and 310 X-ray lines, [15] (150 A in length, 45 A in width) confirmed the very small size and the expected high reactivity of the precipitate.

Propertiesof Plasma Sprayed BioactiveFluorhydroxyapatiteCoatings: X. Ranz et al.

Figure 1. Dissolution curves

457

Figure 2. Weight loss after sonication

Figure 4b represent s the Fourier Transformed infrared spectrum of the precipitate . The poor resolution of the apatitic bands can be related to the poor crystallinity. The bands at 1488, 1424 and 870 cm’^ are due to carbonate ions in labile environment s indicating the high activity of the lay^ [16]. The cartxniate content (about 20% weight) appeared high forcarbonate apatite of type B and corresponde d more probably to sodium-containin g carbonate apatite [18]: Ca7Na2(P04)3(C03)3(OH). In order to evaluate the bioactivity of FHAS , several invitrotests were performed. FHA S coatings appeared to be weakly cyto-toxi c resulting in a partial attachmen t of the human osteoblas t cells : after 6 hours of incubation: 60 % cells attached for HA . and only 35 % for FHA S (Fig. 5). However, FHA S favoured cell proliferatio n during the first 6 days of incubation (Fig. 6). After 27 days, the cell density on FHA S and HA was respectivel y 85 % and 66 % (100 % for the polystyrene test sample). This observatio n could be partly assigned to the effect of fluoride on cell growth [14]. SEM analysis showed that FHA S significantly favoured cell density and dispersion and cyt(9lasmic extensio n in comparison with HA . FHA S coated pins were implanted and compared with HA coated samples in unloaded conditions. Osteointegratio n was assesse d by measuring the percentag e of contact and the percentag e of bone growth. For FHA S and HA bone growth was respectively : 74 % and 80 % after 15 days; 99 % and 98 % after 9 months. This activity could be related to the solubility of the carbonate layer for FHA S and the presence of soluble phase in the HA plasma coating.The quality of the bone/implan t interface was measured by comparing the intrusion torque and extractio n torque of screws 6 weeks after of implantation in loaded conditions [8], both FHA S and HA enhance d the bone/coatin g integration . The histological observation s performed in unloaded and loaded conditions confirmed the osteoconductio n and osteointegratio n properties of the two coatings however a faster degradation of the HA coating was noted. The FHA S coating appeared to be as;

310

Figure 3. SEM inmge of plasma spayed coating after surface treatmen t

Figure 4. a) X-ray diffraction pattern of (rf the pecipitate ; b) FTIR of Uie precipitat e


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Cells /m2 r NOOOO 5000 0

Days l

Figure 5. Cell attachment study

3

6

9

15 2 1 2 7

Figure 6. Cell proliferation study

efficient as the HA coating regarding bone integration but more resistant to successive bone remodelling. CONCLUSIO N Bioactive coatings have to comply with two opposite properties: long-term stability which ensures a good boneAmplant interface and biological activity to favour bone repair. The solution of a bi-layer coating composed of a plasma-sprayed deposit of FHA with good bonding to the metal substrate and a surface lay^ of carbonate âtite ^>pears to be a good alternative although it may be improved. ACKWOLEDGMEN T The present work was a part of a Brite EuRam programme supported by the European Commission.

REFERENCES 1. 2. 3. 4. 5. 6. 7. 8. 9. 11. 12. 14. 15. 16.

Geesink, R., De Groot, K. and Klein, C.P.A.T. /. Bone JL Surg,1988, 70B, 17-22 Klein, C.P.A.T, De Blieck-Hogervorst, J.M.A., Wolke, J.G.C. and De Groot, K. Advances in Biomaterials1990,9.277-282 Van Blitterswijk, C.A., Bowell, Y.P., Rach, J.S., Leenders, H. and Bakker, D. BoneBiomat.Interface1991,295-307 Frayssinet, P., Hardy, D., Hanker, J.S. and Giammara, B.L. Cells and Materials1995,2, 125-138 LeGeros, R.Z., Orly, I., Gregoire, M. and Daculsi, G. Bone-Biomat.Interface1991,76-88 Ranz, X., Gobbi, L., Rustichelli, F., Antolotti, N. and Rey, C. Ceram,Cells and Tissues Volume 2, Ravaglioli 1996, in prepartion Antolotti, N., Chil6, M., Casti, A., Carton, F., Rey, C. and Ranz, X. Patent PR96A000021, 1996 Moroni, A., Orienti, L., Stea, S., Visentin, M. and Nfaltarallo, C. In: BioceramicsVol.8, Pergamon, Oxford 1995,345-350 Mweno, E.C., Kresak, M. and Zahradnik, R.T. Caries Res. 1977,11,142-171 Christoffersen, M.R. and ChristoffCTsen, /. Calcif TissuesInt.1985,37,673-676 W«ig, J., Wolke, J.G.C, Zhang, X. and De Groot, K. In: BioceramicsVol. 8, Pergamon, Oxford 1995,169-175 Farley, J Jl., Wergedal, JE. and BayUnk, DJ. Science1983, 222, 330-337 Sherrer, P. Gdtt.Narch.1918, 2, 98 Vignol, C. Thesis,I.N.P. Toulouse, 1973

Bioceramics, Volume 10 Edited by L. Sedel and C. Rey (Proceedingsof the10th InternationalSymposiumon Ceramics in Medicine, Paris, France, October1997) '1997 Elsevier Science Ltd

LONGER-TER M MECHANICA L AND BIOLOGICA L EVALUATIO N OF TITANIU M ALLO Y COATE D WIT H APATIT E LAYE R W.Q. Yanl*, K. Kawanabel, T. Nakamural, T. Kokubo^ iDept. Of Orthopaedic Surgery, Faculty of Medicine, Kyoto University, Sakyo-Ku, Kyoto, Japan ^Dept. Of Material Chemistry, Faculty of Engineering, Kyoto University, Sakyo-Ku, Kyoto, Japan

ABSTRAC T The objective of this study was the development and use of a biomimetic coating method, which deposits a thin and uniform apatite layer onto titanium (Ti) implants in simulated body fluid (SBF), for improving their bone-bonding ability. The mechanical characteristics and histology of commercially pure Ti and apatite layer-coated Ti-6A1-4V alloy were investigated in rabbit tibiae. Interface failure load was determined using a detaching test after periods of 6,10,24, and 32 weeks. The apatite layer-coated implants exhibited significantly higher failure load than the uncoated control at each time period (all p < 0.001). Histologically, the coated-implants bonded directly to bone via the uniform coating, with no intervening soft tissue. In uncoated controls, there was fibrous tissue intervening at the interface even at longer periods. SEM-EPMA demonstrated a Ca-P-rich layer on the coated implants and direct bonding at the interface. No sign of degradation and delamination of the coating was seen after periods of up to 32 weeks. Our results suggest that apatite layer coating on Ti alloys in situmay improve bone-to-implant bonding and make it as a promising coating material. INTRODUCTIO N Coating of bioactive ceramics on metallic prostheses for promoting direct bone bonding has been a subject of extensive research. The use of HA coatings prepared by plasma-spray methods have shown enhanced quality of bone apposition and excellent early clinical results. However, the integrity of the coating thickness and composition as well as adherence of the coating to metals are still practical problems[l]. There are concerns that the instability of HA-coatings and the generation of HA debris as a result of either degradation or delamination will reduce efficacy in the longer term [2]. We have previously evaluated the surface modification of Ti implants using chemical treatments and reported the formation of bonelike apatite on the treated implants in vitro and in vivo [3,4,5]. The bone-bonding strength of Ti implants was remarkedly increased via the bonelike apatite on the implant surfaces [3]. It is therefore expected that the use of biomimetic process to biologically coat apatite layer on Ti implants in situmay render metals bioactive and may also be the solution to concerns of physically processed coatings. In the present study, the bone-bonding behavior of apatite layer-coated Ti alloys produced by the biomimetic method was evaluated for a longer-term in comparison with commercially pure Ti implants using mechanical test and histological examination. * Present Address: Centre for Biomaterials, University of Toronto, 170 College Street, Toronto, Canada 459

460


MATERIAL S AND METHOD S SBF, the ion concentrations of which were almost identical to those of human blood plasma, were prepared by dissolving NaCl, NaHCOs, KCl, K2HPO4, MgCl2, CaCl, and Na2S04 in ionexchanged and distilled water and buffering at pH 7.40 with Tris buffer. Rectangular plates of Ti-6A1-4V (10 x 10 x 2 mm) were used as substrates. The plates were abraded with #400 diamond paste, cleaned in an ultrasonic bath, and then treated in lOM-NaOH aqueous solution and heated at 600’’C prior to the coating process. Coating with apatite was performed using a biomimetic method. After soaking in SBF with a buffer (pH 7.40) at 36.5 C for 4 weeks, the specimens were removed from the solutions, gently rinsed with distilled water and dried at room temperature. To confirm the deposition of apatite layer on the substrate, the surface of the coated implants were examined by scanning electron microscopy (SEM) (Hitachi S-2500CX), energy-dispersive X-ray microanalysis (EPMA) (Horiba EMAX-3700) and thin film X-ray diffractometry (TF-XRD) (Rigaku, CN2651A). Untreated Ti plates were abraded with #400 diamond paste and used as control implants. All implants were sterilized in ethylene oxide gas before implantation. The plates of apatite layer-coated Ti6A14V alloy (n = 8 for each time point) and Ti control (n = 8 for each time point) were implanted into the metaphyses of mature rabbits for periods of 6, 12, 24 and 32 weeks. Under sterile surgical conditions, a longitudinal incision was made on the anteromedial aspect of the proximal metaphyses. Using a dental burr, a slightly oversized hole was made from the medial to the lateral cortex parallel to the longitudinal axis of the tibia. Then an implant was inserted into the hole by perforating the tibia from the medial to the lateral side. A coated plate was implanted in one leg of the tibia and an uncoated one was implanted contralateral leg as a paired control. At sacrifice, segments of the tibiae containing the implants were excised and subjected to a detaching test to evaluate the bonding strength of the bone-implant interface as reported previously [3,5,7]. The failure load was measured when either the implant became detached from the bone or the bone itself broke. After mechanical testing, all specimens were fixed in a phosphate-buffered formalin solution, dehydrated in serial concentrations of ethanol and embedded in polyester resin. Each large block was sectioned, perpendicular to the longitudinal axis of the tibia, with a diamond band saw (BS30(X), EXAKT, Hamburg, Germany). Sections (5(X) \\mthick) were polished with diamond paper and coated with a layer of carbon for observation using the scanning electron microscope furnished with an energy-dispersive X-ray microanalyzer (SEM-EPMA) and backscatter electron detector. The other sections were further ground to a thickness of 80 |im (Microgrinding MG-4000; EXAKT, Germany) and used for Giemsa surface staining and contact microradiography . RESULT S The SEM-EPMA analyses of the cross-sections of Ti4A16V implants showed that a uniform apatite layer (about 20 |im thick), corresponding to the Ca-P-rich layer, was formed and covered the whole implant surface (Fig. 1). No Ca-P-rich layer was found on the Ti control implants. By TF-XRD, the peaks of apatite were confirmed on the coated implants. In the detaching tests, specimen fracture usually occurred between the bone and the implant, but breakage in the bone was observed in some of the apatite-coated specimens. In uncoated controls, some specimens were separated spontaneously before the test, and hence the failure load was defined as 0 kgf. The failure loads of the coated implants were respectively 1.62–0.68

Longer-TermMechanicaland Biological Evaluationof Ti Alloy Coated With Apatite:W.Q. Yan et al. 461

kgf and 4.13–1.8 4 kgf at 6 and 12 wks, and reached to the failure load of 6.09–2.0 4 kgf at 24 wks and 6.0212.1 0 kgf at 32 wks. In the paired controls, the failure loads were only 0.0310.0 2 kgf and 0.8510. 1 Ikgf at 6 and 12 wks, and increase d to 1.9510.7 6 kgf at 24 wks and 2.0110.8 3 kgf at 32 wks. The effect of the coating treatmen t on the failure load was significant (ANOVA ; p = 0.0001) . Further statistica l testing indicated a significant differenc e in failure load betwee n the coated and uncoate d implants at each time interval (r-test; all p < 0.001) . Histology showed active bone formation on the coated implants and the bone in direct contact with the apatite coating, without any intervenin g soft tissue (Fig. 2). A high degree of bone/implan t contact was demonstrate d for the coated implants by backscatte r electro n microscopy and contact microradiography. In contrast few bone apposition, but many fibrous tissue areas, were observed for the uncoate d implants at the early stages. In the longer-ter m periods, the uniform apatite coating bonded to both bone and substrate with no sign of degradation and delamination . At 32 weeks, the apatite layer was replaced by thin lamellar bone, which remained bonded with the underlying implant. SEM-EPM A showed a uniform apatite layer with a high Ca and P level at the interface betwee n the bone and the coated implants and no Ca-P layer on the uncoate d control at all intervals.

mmm

v\

[Rl^^Qa

tJOeoou) Fig.l SEM-EPM A of the cross-sectio n of the coated implant. A Ca-P-rich layer is deposite d on the Ti alloy surface.

Fig.2. Giemsa surface staining of the coated implant (CI), showing a thin apatite layer bonded to the implant surface and bone (B), with no intervenin g tissue.

DISCUSSIO N Attainment of a more adherent , and stable, coating-implan t interface is considere d to be critical in the longer-ter m implant fixation. In this study, we coated apatite on Ti alloy using a novel biomimetic approach and obtained a thin and uniform coating easily without any complicate d technique . The surface structure measured by TF-XR D and SEM-EPM A indicated bonelike apatite formation and Ca-P-rich layer on the coated implant surface. The mechanica l and biological results demonstrate d that the coated apatite layer had a profound effect on the bone-bondin g behavior of Ti alloy implants. A s previously reported, the formation of a number of apatite nuclei on the substrate with some treatment s can be initiated in simulated body fluids. These apatite nuclei grow

462


spontaneousl y once they are formed on the substrate, and eventuall y a continuous apatite layer is formed by consuming the calcium and phosphate ions in the fluid [4,5,7] . This coating process may be analogous to the process of in vivo apatite formation on bioactive materials, and it appears more bioactive than coatings from conventiona l physical methods which may change the physicochemica l propertie s of apatite due to the high heat involved in those processe s [1,2]. It is well known that chemical bonding betwee n bioactive ceramics and bone is achieve d through a bonelike apatite (Ca-P-rich) layer that formed on the bioactive material surfaces in the body. In the present study, the apatite layer-coate d implants showed direct bone bonding without any fibrous tissue interposition , compared with major areas of fibrous tissue around the uncoated implants. A significantly higher failure load was found in the coated implants than in the uncoate d control at each time period. SEM-EPM A and Giemsa surface staining analyses provided further evidence that the coated implants bonded to bone via the thin coating of bonelike layer with a high Ca and P level. These improvement s could be due to chemical bonding of the bone-implan t interface and strong adherenc e of the coating to the implant. Studies examining the effect of the apatite coating on bone-implan t bonding in the longerterm provided further insight into the nature of the coating/meta l interface . Our finding of the coated implant at the longer-ter m periods showed that the coating adhered to the underlying implant so tightly that fracture occurred within the cortical bone. Apatite layer appeared to be replaced by bony tissue during remodeling without any delamination . This biologic bony substitution might have facilitate d the stability of coating-implan t interface , leading to a longerlasting and strong bone-implan t bonding. Apatite layer-coate d implants produced by the biomimetic method proved to be an effectiv e approach in improving bone-implan t bonding. The use of this method to coat a apatite layer on Ti-based metal indicates that further studies are warranted to investigate the bone-bondin g Ti metal in a load-bearing condition as a more physiological and more durable orthopeadi c implants.

REFERENCES

LFiliaggi, M.J., CoombsJsf.A., Pilliar, R.M., JMiomedMater. Res. 1991,25:1211-1229 . 2.Bloebaum, R.D., Beeks,D., Dorr,L.D., Savory, C.G., Dupont, J.A., Hofmann,A,A., Clin. Orthop.1994,298:19-26 . MaterRes., 1997, 3.Yan, W.Q., Nakamura,T., Kobayashi,M., Kim,H.M., Kokubo,T. JJBiomed. in press. 4. Kokobo,T., Miyaji,F., Kim,H.M., Nakamura,T. J.Am.Ceram.Sic. 1996, 79:1127-1129 . 5.Yan,W.Q., Nakamura,T., Kobayashi,M., Kokubo,T., Kim, H. M., Miyaji, F., Bioceramics 1996, (9): 305-308 . 6.Kitsugi,T., Nakamura,T., Yan, W.Q., Oka,M., Goto,T., Shibuya, T., Kokubo, T., Miyaji, S., J.Biomed.Mater. Res.,1996, 32:149-156 . 7.Yan,W.Q., Kawanabe,K., Nishiguchi,S., Nakamura,T., Kokubo,T. Biomateriah (1997), in press.


ELECTROPHORETI C COATING S OF POROU S APATIT E COMPOSIT E ONT O ALUMIN A CERAMIC S Kimihiro Yamashita, Eiko Yonehara, Jun-ichi Hamagami, Takao Umegaki Department of Industrial Chemistry, Tokyo Metropolitan University 1-1 Minami-Osawa, Hachioji, Tokyo 192-03, Japan

ABSTRAC T The coating of porous hydroxyapatite on alumina and zirconia ceramics was undertaken by the electrophoretic lamination method. The multi-layers were comprised of porous hydroxyapatite, intermediate hydroxyapatite, and adhesive layer of calcium phosphate glass. The open porosity and pore size of the surface layers were adjusted by the addition of graphite or alumina powders. In case of alumina additives, the surface layers were decomposed to tricalcium phosphate during sintering, while hydroxyapatite structure was maintained in graphite-added surfaces. K E Y W O R D S : Hydroxyapatite, Porous Apatite, Electrophoretic Coatings, Alumina Ceramics, Composite Coating, Bonelike Crystal Growth INTRODUCTIO N Electrophoretic deposition (EPD) combined with sintering is reportedly one of the practical coating methods of hydroxyapatite (Caio(P04)6(OH)2, HAp) [1-5]. We have carried out EPD coatings of apatite on the ceramics of alumina and yttria-stabilized zirconia for the biomedical use [5]. For this purpose, porous surface layers with HAp structure are desirable. The strong adhesion of the coated layers to substrates is cdso required. To achieve such complex aim, we carried out the EPD lamination of the porous surface, intermediate HAp, and adhesive layers on alumina ceramics. In the present EPD method, the adhesion layers to ceramic substrates were first deposited by the EPD of the mixed powders of HAp and calcium phosphate glasses (CaQP205, CP) with a lower melting point (1(XX)-12(X)**C) onto ceramic alumina substrates according to the previous work [4]. Then HAp intermediate layer was formed on the first layer. The porous surface layers were lastly formed by the co-deposition of (1) HAp and alumina powders or (2) HAp and graphite powders; the pores were generated during sintering. This paper mainly reports the EPD preparation of porous surfaces as well as the lamination procedure. MATERIAL S AN D METHO D Starting powders of HAp were prepared by a wet chemical method. CP powders for adhesion layers were obtained by crushing the dried agglomerates which were obtained by the reaction of CaC03 with phosphoric acid. The reactant powders had the atomic ratio of Ca/P=l/2. Because CP powders were scarcely deposited, they were co-deposited with HAp. The surface layers were formed by the co-deposition of HAp with alumina (40wt%) or graphite powders (lwt%). Commercial alumina powders with an average diameter (d)=ljim were used, and graphite powders with d=10 and 25jim were prepared by scratching, crushing and sieving from the bulk graphite. 463

464


The EPD was undertaken by applying dc voltage of lOOV/cm for several min to those mixed or single powders suspended in a mixed solvent of acetylacetone and ethanol (Acac/EtOH). After drying, the multi-layered alumina substrates were sintered at ISOOT. The coated aluminas were characterized by XRD and IR analyses. In order to evaluate the effectiveness of porous surface, the coated specimens were immersed in a simulated body fluid (SBF) for a week. Scanning electron microscopy was undertaken for the observation of bone-like crystal growth at deep parts in a coated layer. RESULT S AN D DISCUSSIO N Figure 1 shows the laminated layers (A) consisting of adhesion (B), intermediate (C), and porous surface (D). The adhesive and intermediate layers were well densified by sintering. The adhesion layer, which was composed of HAp and CP, was reacted to calcium phosphate compound with a lower melting point (-1100**C). The adhesion strength was lOMPa, which was influenced by the surface roughness. Figure 1 also indicates that the surface layer was porous. The result was explained by the inhibition of densification of HAp powders by the reaction between HAp and alumina, which the XRD analysis attributed to the decomposition of HAp to tricalcium phosphate (TCP) phase and the formation of CaAl204. The pore size as well as porosity was dependent on the quantity of added alumina powders, however, the decomposition was accelerated.

Porous Surface Layer HA p Layer Adhesive CP Layer Ceramic Substrate Figure 1. A schematic view of EPD laminated apatite composite layers on a ceramic substrate (A), and the scanning electron micrographs of sintered layers (porous surface (B), intermediate HAp (C), and adhesive CP (D)). The each bar indicates lOjim.

ElectrophoreticCoatings of Porous Apatite Compositeonto Alumina Ceramics: K. Yamashitaet al.

465

In relation to bioactivity, the decomposition of HAp phase is unfavorable. When the addition of alumina was reduced to 10wt%, some portion of HAp was maintained. We also created pores by the addition of graphite powders to HAp. The mixed powders of C and HAp were well deposited by EPD in Acac/EtOH solvent. Figure 2 shows SEM photographs of the sintered surface of C-added HAp, indicating the effect of the size of graphite powders on the production of porous surfaces. The pores were created by burning of graphite during sintering. Both pore size and porosity were dependent on the amount and size of added graphite powders. Figures 3 (A)-(P^ compares the effectiveness of porous surfaces for the crystal growth in SBF; whereas the crystal growth rarely took place at grains 20jim deeper from the surface in non-added HAp layer (C), bonelike crystals were observed even at the part 80m deeper from the surface at porous HAp layer {¥). This result is attributed to the easy immersion of SBF solution into a porous layer. SUMMAR Y The effectiveness of the electrophoretic deposition technique was shown for the composite coating of porous hydroxyapatite surface and adhesive layers on ceramic substrates. The creation of pores at a surface was carried out by the addition of alumina or graphite powders to hyckoxyapatite.

Figure 2. The scanning electron micrographs of non-added HAp (A), and porous HAp surfaces (B-D), which were sintered at 1300C. The pores of the surfaces were adjusted by the addition of lOmg graphite powders with the diameter (d)=10fim (B) and 25\km (C), and 20mg with d=25^m (D).

466


Figure 3. SEM photographs of fractured surfaces of non-added HAp (A, B, C) and porous HAp (D, E, F) layers, which were immersed in advance in a SBF. The photographs were taken at the points from 10 (B), 20 (C, E), and SOîm (F) from the surfaces, respectively.

REFERENCES

1. Ducheyne. P., Radine S., Heughebaert M., Heughebaert J.C., Biomaterials,1990, 11, 244254. 2. Ducheyne P., van Raemdonc W., Heughebaert J.C, Heughebaert M., Biomaterials,1986, 7, 97-103. 3. Umegaki T., Hisano Y., Yamashita K., Kanazawa T., Gypsum & Lime, 1989, No. 218, 2428. 4. Ding X., Yamashita K., Umegaki T., J .Ceram Soc.Jpn., 1995, 103, 867-869. 5. Nagai M., Yamashita K., Umegaki T., Phos, Res. Bull, 1991, 1, 167-172. 6. Garvie R.C., Urbani C , Kennedy D.R., McNeuer J.C, J. Mat, ScL, 1984, 19, 3224-3228. 7. Christel P., Meunier A., Heller M., Torre J.R, Peille CN., 7. Biomed.Mat Res,, 1989, 23, 45-61. 8. Hayashi K., Matsuguchi N., Uenoyama K., Kanemaru T., Sugioka Y., 7. Biomed,Mat, Res,, 1989, 23, 1247-1259.


OSSEOINTEGRATIO

N IN EXPERIMENTA L HA-COATE D FEMORA L STEM S

E. De Santis, G. Rinonapoli, C. Doria, A. Manunta, MC. Sbemardori University of Sassari, Italy, Orthopaedic Dept, C.P. 40, 07100 Sassari, Italy.

ABSTRAC T The purpose of the study is to evaluate the bone-ingrowth in experimental implants of HA-coated femoral stems. Twelve sheeps were submitted to hip hemiarthroplasty with a specially designed femoral component HA-coated in the proximal 2/3 of the stem. The animals were killed at regular intervals of time after surgery. Both femora have been removed, fixed in neutral buffered formalin and submitted to plain radiographs, CT-scan, DEXA and Scanning electron microscopy. The osseointegration around the implants occurred in eleven specimens, one stem was loosened by 30 days. At 15-30 days the apposition of woven inunature bone was evident, and mature lamellar bone at 45 days. At this time, the gap between the endosteum and the HA-coated surface wasfilledby bridges of lamellar bone; this bone-ingrowth was prevalent in the proximity of the grooves and on the posterior surface increasing with time, with marked condensation of new formed bone, widely remodeled trabeculae. The layer of HA-coating tends with time to resorb; it is prevalently a slow resorption, with gradual replacement of HA by new formed bone confirming its osteoconductive properties. KEYWORDS : Hydroxyapatite, Hip arthroplasty. Prosthesis, Sheep, Osseointegration, Osteoconduction INTRODUCTIO N The goal of cementless prosthetic implants is to obtain an adequate biological stability by bone ingrowth. For several years the aim was to find the material which could allow and promote an optimal osseointegration without biological damage. Hydroxyapatite (HA), with its properties of atoxicity, antigenic inactivity, no carcinogenicity, bioactivity and biocompatibility, is the most studied material in this field. The large amount of studies on the osteoconductive property of this ceramic have reported good results [1,2,3,4,5,6,7]. We investigated about this topic with some experimental studies [8], which widely confirmed what had been found by other Authors. In the present experimental study we investigated the process of osseointegration in HA-coated femoral stems. MATERIA L AND METHOD S Twelve 2-3 years old sheeps, weighing 35-38 Kg., were submitted to unilateral hip hemiarthroplasty, with a specially designed femoral component. The stem, made of a special 467

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titanium alloy with a low modulus of elasticit y (Tilastan), was straight and cylindric in design, with a cx)llar and two longitudinal deepe r grooves in the diaphyseal portion and 11 transverse grooves each side. The proximal 2/3 of the stem had a circumferentia l HA-plasma sprayed porous-coatin g (thickness 105-10 8 micron, purity 95%, cristallinity 70%, porosity 11%) with 50-110 \i pores. A 22 mm Co-Cr alloy head was used, with a 10-12 nun morsetaper. Under endotrachea l general anaesthesia , an anterolatera l approach to the hip joint was performed. A combination of streptomyci n and penicilline (1.200.00 0 units) was administere d subcutaneousl y for 7 days postoperatively . The animals were allowed imrestricte d weight-bearin g immediatel y after surgery and walked daily during the whole postoperativ e period. They were killed by a barbiturate overdose at regular intervals of time (15, 30, 45, 60, 90, 120, 180, 200, 270, 360, 540 days). Both femora and the pelvis were harveste d and the soft tissue removed. High definition radiographs (AP+LL) , CT-scan and DEX A were taken of the implants. All specimen s were fixed in 10% buffered formaline. Cross-section s were obtained by a high-spee d diamond saw in the proximal femur at 4 levels: intertrochanteric , upper, middle and distal portion of the diaphysis filled by the stem. The sections , prepared for the observatio n by the classic procedure of fixation in 2.5% glutaraldheyde , post-fixatio n in 1% Osmium tetroxide , dehydration in acetone , drying in the critical point in C02 and metallize d in gold, were observed under scarming electro n microscope (ZeissDSM962). RESULT S In the early phases (15, 30, 45 days) the HA-coating maintained its original thickness and appeared strongly adherent to the metal surface. The gap, present at HA-bone interface , was progressivel y filled by new formed bone. Apposition of woven inunature bone, was evident at 2-3 weeks, followed by lamellar new bone with a marked trabecular orientatio n at 4-6 weeks. The bone ingrowth occurred in proximal and middle levels of the stem, where HA-coated surface was present. The process was greater in the posterior part of the bone-implan t interface . The new formed trabeculae formed bridges in the gap betwee n the endosteu m and the HA-coated surface, especiall y in the proximity of the transverse grooves and close to the medial cortex (zone 7 by Gruen). The entity of the bone ingrowth increase d with time and we saw, at 180-20 0 days, at the metaphysea l level, a great condensatio n of widely remodele d bone trabeculae which appeared to become thicker, as shown by X-ray and CT-scan. In the diaphyseal HA-coated portion of the stem, we observed the same stages of the process of osseointegratio n that we had seen in the proximal metaphysea l portion, but were slower in their development . At 200 days, it was possible to detect a different aspect betwee n the compact host bone and the new-forme d lamellar bone, that appeared emiched by residuals of amorphous HA crystals. Wefrequently detecte d an evident line of separation betwee n the host bone and the new-forme d bone, and this is correlate d to the different patterns of bone that present different mechanica l properties . In the distal uncoate d portion, the implants were surrounded by intervenin g fibrous tissue that was transformed into lamellar bone at 270 days. The beginning and the evolution of the osteogeneti c activity at the bone-implan t interface was associate d with evident changes of the HA . An actual phenomeno n was the progressive reduction of the thickness of the bioactive coating, until its complete disappearance , which was observed at around 200 days from surgery (Fig. 1). The gap.

Osseointegrationin ExperimentalHA-Coated Femoral Stems: E. De Santis et al.

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previously filled by the HA , was then replaced by lamellar bone, differentl y arranged in the different sections of the femur. The process of gradual replacemen t of the HA by bone was precede d by morphological and structural changes, already observable at 60 days, consisting of lacunae and cracks with varying orientation . At greater distance of time (90 days), the confluenc e of crack lines led to the microfragmentatio n of the HA . In some sections we observed, already by 60 days, the presence of coarse sods in the point of maximum contact betwee n bone and stem. After resorption had been complete d (270 days) it was possible to detect amorphous residuals of synthetic HA mixed with new-forme d bone. The latter bone was totally different from the host bone and fi-omthe bone formed in the early phases of the experiment . It was lamellar bone, enriched by residuals of HA and showed a very peculiar arrangement: the lamellae, not yet mechanicall y solecitated , had not the same regular arrangement of the host bone. The resorption of the HA layer led to the appearance of a gap betwee n the new bonefi-ontand the bare metal. This gap underwen t a progressive filling with time, simultaneousl y with the forward movemen t of the new bone layer, as observed at 270 and 360 days (Fig. 2). The adaptive bone remodeling led to thickening with increase d density at DEX A in the medial and anterior cortex and thinning in the lateral and posterior cortex; it started at 45 days and progressed with time, until 540 days. Spongiosizatio n was evident, beginning at 90 days, in zone 7 under the medial collar and anteriorly, in zone 8. It increase d until 540 days. Spongious hypertrophy in zone 1 appeared at 30 days, progressing imtil 90 days and decreasing after 180 days, as shown by CT-scan and DEXA ; this technique show also an early demineralizatio n of the great trochante r that progress until 90 days and decrease progressivel y so than after 540 the mineral density at this level appear normal.

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40. 3. Tani, N. JapaneseJ. Medic. Instrument,1988,58,266-273. 4. Tanihara, M. etal.,JapaneseJ. Artific.Organs, 1989,18,15-18. 5. Agishi, T. etai, JapaneseJ. Artific.Organs. 1991,20,318-323. 6. Takashima, S., etal..In: Bioceramics.,Volume 9, Pergamon, Oxford, 217-220. 7. Kitano, Y. etai, JapaneseJ. Artific.Organs, 1988,20,1491-1496.


PREPARATIO N OF P^-IMPLANTE D YjOj-AljOj-SiO j GLAS S FOR RADIOTHERAP Y OF CANCE R M. Kawashita\ F. Miyaji\ T. Kokubo\ G. H. Takaoka^ I. Yamada^ Y. Suzuki^ and M. Inoue^ ^Department of Material Chemistry, Faculty of Engineering, Kyoto University, Sakyo-ku, Kyoto 606-01, Japan, ^lon Beam Engineering Experimental Laboratory, Faculty of Engineering, Kyoto University, Sakyo-ku, Kyoto 606-01, Japan, ^lon Engineering Research Institute Corporation, Hirakata, Osaka 573-01, Japan ABSTRAC T P^ ion was implanted into a Y203-Al203-Si02 glass, which has already been used for radiotherapy, at 200 keV with a dose of 1x10^^ cm’^ in order to increase its treatment effect. The phosphorus was distributed up to the glass surface and a part of it near the surface was oxidized. The P^-implanted Y203-Al203-Si02 glass released appreciable amounts of P, Y and Si into water at 95 C for 7 d, although the P^-implanted silica glass hardly released these elements under the same condition. This result indicates that P^ ion must be implanted with lower doses than Ix 10^^ cm"^ or at higher implantation energies than 200 keV in order to obtain highly chemically durable P-containing Y203-Al203-Si02 glass, preventing the distribution of phosphorus to the surface layer. KEYWORDS : Phosphorus, Ion implantation, Y203-Al203-Si02 glass. Radiotherapy INTRODUCTIO N Radiotherapy is one of the effective treatments of cancers. External irradiation, however, often causes damages to healthy tissues. It has been reported that a 17Y203-19Al203-64Si02 (mol%) glass is usefiil for in situirradiation of cancers [1]. Y-89 in the glass can be activated to P-emitter Y-90 with 64.1 h half-life by neutron bombardment. Microspheres of the activated glass can give large local irradiation of the short-ranged highly ionizing p-ray to the tumors with little radiation dose of neighboring organs, when they are injected to the tumors. Y-90, however, may result in the substantial decay before the treatment owing to short half-life of 64.1 h. P-31 with 100% natural abundance similar to Y-89 can be activated to pemitter P-32 with 14.3 d half-life by neutron bombardment. But highly phosphorus-containing glasses prepared by the conventional melting method are usually less chemically durable. It can be expected that a glass, which is more effective for radiotherapy, can be obtained by P^ ion implantation into Y203-Al203-Si02 glass. The present authors previously showed that P^ ion can be successfiilly implanted into a silica glass under 200 keV with a dose as high as 1x10^^ cm"^ without giving adverse effect on its high chemical durability [2-4], In the present study, P^ ion was implanted into 17Y203-19Al203-64Si02 (mol%) glass with a dose of 1x10^^ cm"^ under 200 keV in order to obtain the glass with higher treatment effect of radiotherapy. The state of the implanted phosphorus, and structural change of the glass surface due to ion implantation were examined. Chemical durability of the P^-implanted 531

532 Bioceramics Volume10

glass was discussed in terms of its surface structure. MATERIAL S AND METHOD S A glass batch of 17Y203-19Al203-64Si02 (mol%) composition consisting of reagent-grade chemicals of Y2O3, AI2O3 and Si02 was melted in a platinum crucible at 1600 C for 2 h, poured on a stainless steel plate to be formed into plates 1 mm thick. The obtained glass was cut into rectangular specimens 10x10x1 mm^ in size, polished with 3-4 |im diamond paste, washed with pure acetone in an ultrasonic cleaner. The glass plate was annealed at 850 C for 1 h in order to eliminate the strains. The glass was implanted with P^ ion at 200 keV with a dose of 1x10^^ cm’^. The distribution of phosphorus was measured by a Rutherford backscattering spectrometry (RBS) using 2 MeV "^He^ ions with 170 incident angle. The state of phosphorus in the glass was investigated by measuring survey spectrum with an X-ray photoelectron spectroscope (XPS) (MT-5500, ULVAC-PHI Co. Ltd., Chigasaki, Japan), using MgKa X-ray as the source. The P^-implanted glass was soaked in 20 ml of distilled water at 95 C for 7 d in a polypropylene bottle, shaken at a rate of 120 strokesmin’^ with a stroke length of 3 cm. The concentrations of the phosphorus and silicon released from the glass were measured by an inductively coupled plasma atomic emission spectrometer (SPS-1500 VR, Seiko Instruments Inc., Tokyo, Japan). RESULT S AND DISCUSSIO N Figure 1 shows the RBS spectrum of P^-implanted Y203"Al203-Si02 glass, in comparison with that of unimplanted original glass. Judging from a broad peak at 200-300 in channel number, it is assiuned that P^ ion was successfully implanted into the glass, although the peak of phosphorus overlapped those of aluminum and silicon. A rising was observed at about 400 in channel number for PMmplanted Y203-Al203-Si02 glass. This indicates that yttrium ions moved to the glass surface during the ion implantation. Figure 2 shows the P2p XPS spectra of P^-implanted Y203-Al203-Si02 glass. Two peaks.

P’^-implanted

Unimplanted Y

1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 1

200

300

400

Channel nunnbe r

1

500

Figure 1 RBS spectrum of P^-implanted Y203-Al203-Si02 glass, in comparison with that of unimplanted original glass.

134 130 Binding energy /eV

Figure 2 P2p XPS spectra of P^-implanted Y203-Al203-Si02 glass.

P^-Implanted Y2O3 -AI2O3 -Si02 Glass for Radiotherapyof Cancer: M. Kawashita et al.

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assigned to elementa l phosphorus (130 eV) and oxidized phosphorus (134 eV), were observed at the surface. With increasing depth, the intensity of the peak ascribed to the elementa l phosphorus increase d while that ascribed to the oxidized phosphorus decreased . This means that a part of phosphorus exists as oxidized one near the surface, although most of it exists as elementa l colloids in the glass. Figure 3 shows the concentration s of phosphorus, silicon and yttrium release d from the PMmplanted Y203-Al203-Si0 2 glass into water at 95 C for 7 d. The glass release d appreciable amounts of P, Y and Si into the hot water. The appreciable release of yttrium and phosphorus may be attribute d to the surface localizatio n of yttrium (Fig. 1) and the formation of chemically less durable phosphorus oxide near the glass surface (Fig. 2), respectively . The release s of larger amount of the yttrium and the phosphorus might enhance the release of silicon. Figure 4 shows the RBS spectra of PMmplanted Y203-Al203-Si0 2 glasses before and after soaking in water at 95 C for 7 d. Peak area ascribed to phosphorus remarkably decrease d after the soaking in hot water. This result indicates that the implanted phosphorus was release d into water by the soaking, which is consisten t with the result shown in Fig. 3. For silica glass implanted with P^ ion at 200 keV with a dose of 1x10^^ cm’^ [4], the implanted phosphorus was localized only in deepe r regions, and hence the glass showed high chemical durability. For P^-implanted Y203-Al203-Si0 2 glass, the implanted phosphorus can not penetrat e into the deep region of the glass, because it contains heavy Y-89. Consequently , the phosphorus, which was implanted into Y203-Al203-Si0 2 glass, was oxidized to form chemically less durable phosphorus oxide at the glass surface and the glass showed lower chemical durability compared with P^-implanted silica glass. This means that the implanted phosphorus must be localized in deep region for keeping high chemical durability of the glass even after the P^ ion implantation . It is therefor e suggeste d that P^ ion must be implanted with lower doses than Ix 10^^ cm"^ or at higher implantation energies than 200 keV in order to obtain highly chemically durable P-containing Y203-Al203-Si0 2 glass, preventin g the accumulatio n of phosphorus in the surface layer.

Unim planted ^ P -implante d Unlmplante d d ^ ^ ^^ P’^-implante Unimplante d ^ P’^-implante d 0

0.5

1

1.5

Concentratio n /pp m

2

Figure 3 Concentration s of P, Y and Si release d from P^-implanted Y203-Al203-Si0 2 glass soaked in water at 95 C for 7 d.

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200

300

400

Channel number

500

Figure 4 RBS spectra of P^-implanted Y203-Al203-Si02 glasses before and after soaking in water at 95 C for 7 d. CONCLUSION S Y203-Al203-Si02 glass was implanted with P^ ion at 200 keV with a dose of 1x10^^ cm ^ The P^-implanted Y203-Al203-Si02 glass released appreciable amounts of P, Y and Si, although the P^-implanted silica glass hardly released these elements under the same condition. This is because the implanted phosphorus was widely distributed up to the surface and a part of it near the surface was oxidized in the former glass, while it was localized in deeper region in the latter glass. P^ ion must be implanted with lower doses than 1x10^^ cm"^ or at higher implantation energies than 200 keV in order to obtain highly chemically durable P-containing Y2O3-AI2O3Si02 glass, preventing the distribution of phosphorus to the surface layer. ACKNOWLEDGMENT S We thank Radiation Laboratory of Nuclear Engineering, Kyoto University, for RBS measurement. This work was supported by Grant-in-Aid for Young Scientists from Ministry of Education, Science, Sports and Culture, Japan.

REFERENCES L

Ehrhardt, G.J. and Day, D.E., NucL Med Biol.1987,14, 233-242.

2. Kawashita, M., Miyaji, F., Kokubo, T., Takaoka, G.H. and Yamada, I., J.Ceram.Soc, Jpn. 1996, 104, 710-714. Kawashita, M., Miyaji, F., Kokubo, T., Takaoka, G.H., Yamada, I., Suzuki, Y. and Kajiyama, K., Nucl Inst Meth.in Phys. Res. B, 1997, 121, 323-327. Kawashita, M., Miyaji, F., Kokubo, T., Takaoka, G.H. and Yamada, I., In: Proc. 2ndlntl Meet.PacificRim Ceram. Soc, The Austalasian Ceram. Soc., 1996, in press.

Bioceramics, Volume 10 Edited by L, Sedel and C. Rey (Proceedings of the 10th InternationalSymposium on Ceramics in Medicine, Paris, France, October 1997) '1997 Elsevier Science Ltd

NEW FERROMAGNETIC BONE CEMENT FOR LOCAL HYPERTHERMIA K. Takegami*, T. Sano*, H. Wakabayashi*, J. Sonoda*, T. Yamazaki*, S. Morita**, T. Shibuya**, A. Uchida* *Department of Orthopedic Surgery, Mie University Faculty of Medicine, Edobashi 2-174, Tsu, Mie 514, Japan **Nippon Electric Glass Co., Ltd. ABSTRAC T We have developed a ferromagnetic bone cement as thermoseed to generate heat by hysteresis loss. This material resembles bioactive bone cement in composition, with a portion of the bioactive glass ceramic component replaced by magnetite powder. The temperature of this thermoseed rises in proportion to the weight ratio of magnetite powder, the volume of the thermoseed, and the intensity of the magnetic field. The heat-generating ability of this thermoseed implanted into rabbit tibiae was investigated by applying an alternate magnetic field In this system, it is very easy to increase the temperature of the thermoseed in bone beyond 50X1 by adjusting the above-mentioned control factors. Localized hyperthermia in an experimental bone tumor model was induced and inhibition of the tumor growth by this hyperthermic therapy was confirmed radiologically and histologically. These results demonstrate that ferromagnetic bone cement is useful for the treatment of musculoskeletal tumors. KE Y WORD S ferromagnetic thermoseed, hyperthermia, bone cement INTRODUCTIO N The efficacy of hyperthermia for tumors heated above 42X has been confirmed by many researchers [1,2]. In the deep region, however, it is difficult to heat the tumor selectively. As a result of this problem, various methods have been developed for localized hyperthermia by magnetic induction heating [3,4]. We have developed a new ferromagnetic thermoseed of the cement type for local hyperthermia of the tumors in the deep regions, especially in bone. This material generates heat by applying an alternate magnetic field. It resembles bioactive bone cement [5] in composition, with a portion of the bioactive glass ceramic component replaced by magnetite (Fe304) powder and thus can be molded into various shapes. In this stucfy, we investigated this cement’s heat-generating ability under various conditions for its efficacy in an experimental model of bone tumors. MATERIAL S AN D METHOD S Ferromagnetic bone cement was supplied by Nippon Electric Glass Co., Ltd.( Ohtsu, Japan). This cement was composed of the two-paste type. In one paste, benzoyl peroxide (0.1 wt%) was dissolved and, in the other paste, N, N-dimethyl-p-toluidine (0.1 wt% ) was dissolved. The paste consisted of magnerite ( Fe304) and silica glass (SiOj) powders as filler and a BIS-GMA-based resin. The ratio of the filler to the resin was 9:1 (weight: weight). The average particle size of the magnetite powders was 535

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13 |Lim and that of the siHca glass powder 3 |Lim. The BIS-GMA- based resin was composed of bis-a-glycidylmethacrylate and triethylene-glycol dimethacrylate. Two pastes of the same weight were mixed and kneaded by hand for 1 minute. The mixture was molded into various shapes andsohdified in about? minutes. There was little rise in temperature of the cement during polymerizing. Investigation of heat-generatin g ability under various conditions The heat-generating ability of the ferromagnetic bone cement was investigated as a function of the content of magnetite, the volume of the cement, and the intensity of an appHed magnetic field. Four kinds of cement (containing this powder at 10%, 20%, 40%, or 80% weight ratio) were molded into a cubic block of 20 x 20 x 20 mm for studying magnetite content. The cement-containing magnetite at 40% weight ratio was molded into rectangular blocks of 20 x 20 x 20 mm, 20 x 20 x 10 mm, and 20 X 20 X 5 mm relative to cement volume. Each of these blocks was subjected to an alternate magnetic field to generate heat. The rectangular block (20 x 20 x 5 mm) was prepared relative to the intensity of the applied magnetic field and then subjected to altemate magnetic fields of various intensities. After the blocks were heated for ten minutes, the temperature of the surface on the blocks was measured with a fluoropdc thermometer (model 3000; Laxtron, Mountain View, CA, U.S.A.). Investigation of distribution of heat in the bone Japanese white rabbits weighing 2.0 to 2.5 kg were anesthetized with 25 mg/kg intravenous pentobarbital sodium. The cement (containing magnedteat 50% weight ratio), a molded pillar with a diameter of 4 mm and a length of 25 mm, was inserted into the medullary canal of the rabbit tibiae. After closure of the wound, the lower leg containing the cement was placed in an altemate magnetic field and heated Five sensors from the thermometer were used for thermometry in the leg. In the portion containing the implanted cement, one of the sensors was inserted at the surface of the medullary canal and two were at the medial and lateral interfaces between bone and muscle. In the distal portion, one was at the medullary canal at a distance 10 mm from the cement and one was at medial interface between bone and muscle at a distance 9 mm from the cement (Fig. 1 A). The temperature of the cement was adjusted to 50- 6 0 ^ . Local hyperthermia for experimenta l bone tumor Japanese white rabbits weighing 2.0 to 2.5 kg were implanted with VX2 tumor blocks, 2 x 2 x 2 mm, in the right tibiae under local anesthesia. Two weeks later, a second operation was performed. At this time, the rabbits were divided into two groups: the no treatment group and the hyperthermia

d e 16 24 32 4 0 time (min.) Figure 1 A: the position of thermometry. Surface on the cement (a), medial (b) and lateral (c) interfaces between bone and muscle, cortical surface(d) and medullary canal (e) in distal portion. B: time/temperature curve.

B

0

8

New FerromagneticBone Cementfor Local Hyperthermia:K.^Takegamiet al.

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group. In the hyperthermia group, each rabbit was anesthetized with intravenous pentobarbital sodium. The medullary canal of the right tibia was curretted and the bone defect was filled with the ferromagnetic cement. An alternate magnetic field was applied to the right leg for 50 minutes. During heating, the temperature on the cortical surface of the tibia was maintained at 42-45"C. All of the rabbits were killed 5 weeks after transplantation of the VX2 tumor and radiographs were taken of the right leg. The right tibia was removed and evaluated histologically. RESULT S AN D DISCUSSIO N Heat-generatin g ability under various conditions An increase of magnetite content from 20% to 80% caused an increase in temperature of the material (Fig. 2a). Maximum temperature of 80% magnetite was 6IX! in the magnetic field of 80-Oe, 100 kHz. In regard to the cement volume, the temperature of the cement rose at afixedrate in proportion to the volume (Fig. 2b). The temperature of the cement kept rising as the intensity of the alternate magnetic field was increased (Fig. 2c). These results were similar to other ferromagnetic thermoseeds that produced heat by hysteresis loss [6]. As this material is of the cement type, an accurate estimate of the volume of the implanted cement can be determined prior to implantation. Therefore, it is important that the intensity of the magneticfieldcan be easily adjusted by the power output of the AC magnetic field generator and that the temperature of the cement also can be controlled. Distribution of heat in the bone In the rabbit tibiae, the cement was heated to 50X! within four minutes and then the temperature of the cement was maintained within a range of 50-60X! for approximately one hour. The temperature at the interface between bone and muscle increased to 43-45X and was maintained for the duration of this experiment (Fig. IB). Cell destruction specifically occurred in tumor cells without any apparent damage to normal cells at the latter temperature [ 1 ]. This demonstrated that heat damage to soft tissue around the bone is prevented by this method. On the other hand, the tissue surrounding the cement was heated by conduction and the temperature of the bone adjacent to the cement rose easily above 45X1. A temperature above 45X! is better for obtaining rehable tumoricidal effects to any tumor cells remaining after surgical resection. However, in the distal portion implanted with the cement, the temperature at the cortical surface did not increase above 42X! and the temperature at the medullary canal was maintained at 37X!. At some distance from the thermoseed, however, bone marrow was insufficientiy heated for hyperthermia. Local hyperthermia for experimenta l bone tumor In all rabbits, the radiolucency within the area of implantation was seen on radiographs taken 2 weeks post tumor transplantation.

0 5 10 15 20 60 10 0 14 0 18 0 20 40 60 b height of block (mm) content (96) Figure 2 The temperature change of this material a: relationship between temperature and volume, (base of block, 20 x 20 mm. Content of magnetite, 40%. Magnetic field, 70 Oe.) b: relationship between temperature and content of magnetite, (block, 20 X 20 X 20 mm. Magnetic field, 110 Oe.) c: relationship between temperature and intensity of magnetic field, (block, 20 x 20 x 5 mm. Content of magnetite, 40%.)

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Figure 3 Radiological findings of the no treatment group (A) and the hypertheraiia group (B), Pathological finding of the hyperthermia group (C). Figure 3 A andB show radiological findings of each group 5 weeks after transplantation. Although marked destruction of the tibia was seen and pathologic fractures occurred in the no treatment group, radiolucency, cortical destruction, and pathologic fractures were not seen in the hyperthermia group. Pathologic evaluation revealed viable peripheral and necrotic central portions in the no treatment group. In the hyperthermia group, necrotic portions expanded from the cement (Fig. 3C). The hind legs of rabbits were temporarily swollen after heat production, but no irreversible complications, such as skin necrosis, occurred These results demonstrate that the bone tumor was sufficiently and selectively heated in this system. CONCLUSIO N This study demonstrated that ferromagnetic bone cement is useful for local hyperthermia in bone. REFERENC E [1] Overgaard, J. Cancer, 1977, 39, 2637-2646. [2] Dichson, J.A. and Ellis, H.A. Nature, 1974, 248, 354-358. [3] Meijer, J. G., van Wieringen,N., Koedooder,C., Nieuwenhuys,G.J. and van Diji, J.D.P. Med. Phys. 1995, 22, 101-104. [4] Ohura, K., Kenaga, M., Nakamura, T., Yamamuro, T., Ebisawa, Y., Kokubo, T., Kotoura, Y. and Oka, M. J. Applied Biomaterials, 1991, 2, 153-159. [5] Kawanabe, K., Tamura, J., Yamamuro, T., Nakamura, T., Kokubo, T. and Yoshihara, S. J. Applied Biomaterials, 1993, 4, 135-141. [6] Jordan, A., Wust, P., Fahling, H., John, W. and Hinz, A., Felix, R. INT. J. Hyperthermia, 1993, 9, 51-68.

BIOCERAMICS PROCESSING



ADSORPTIO N OF L- LYSIN E ONT O SILIC A GLASS : A SYNERGISTI C APPROAC H COMBININ G MOLECULA R MODELIN G WIT H EXPERIMENTA L ANALYSI S Robert A. Latour Jr.^, Jon K. West^, Larry L. Hench^, Sharon D. Trembley ^, Yuan Tian^, Gary C. Lickfield"^, Alfred P. Wheeler^ 1 Department of Bioengineering, Clemson University, Clemson, SC, USA 29634 2 Department of Materials Sci. & Engineering, University of Florida, Gainesville, FL, USA 32611 3 Department of Materials, Imperial College, University of London, London, UK SW7-2BP 4 School of Textiles, Fiber and Polymer Science, Clemson University, Clemson, SC, USA 29634 5 Department of Biological Sciences, Clemson University, Clemson, SC, USA 29634 ABSTRAC T As a basic science approach to the problem of protein adsorption on implant surfaces, experimental and molecular modeling studies were conducted to study the adsorption of a protein residue (lysine) onto silica glass. Experimentally, 4 molecular weights of poly-L-lysine were adsorbed onto glass microspheres at 4 temperature levels. Adsorption enthalpy for each molecular weight was determined from the adsorption isotherms and plotted versus the degree of polymerization to estimate the enthalpy per adsorbed residue. Molecular modeling was also performed using a CAChe Worksystem. The geometries of the silica rings and lysine were optimized using the AMI semi-empirical method and the secondary structure of poly-L-lysine was evaluated using the MM2 molecular mechanics method. The modeling results yielded reasonably close correlation with the experimentally measured adsorption energy (-2.6 kcal/mol), and provided insights into likely adsorption mechanisms. KEYWORDS : adsorption, lysine, proteins, molecular modeling, computational chemistry INTRODUCTIO N Although numerous investigations have been conducted to investigate the adsorption of whole proteins to surfaces [1-5], because of the large numbers of variables involved, these studies have yielded relatively little information regarding protein conformational changes following adsorption. It is proposed that alternative approaches to protein adsorption are necessary to develop an understanding of protein adsorption behavior. Proteins are complex macromolecules with a minimum of 3 levels of hierarchical organization termed primary, secondary, and tertiary structure. It is proposed that to begin to understand, and thus predict the behavior of adsorbed proteins, we must first develop an understanding of the adsorption process at the most fundamental level (i.e. primary structural level). This understanding should then provide a basis to begin to sequentially address the adsorption process involving the higher levels of structure. This paper presents the results of a study of the adsorption behavior of proteins at the primary structural level. A combined experimental [6] and molecular modeling approach was taken which first involved the analysis of adsorption isotherm data to experimentally determine the adsorption enthalpy for mid-chain amino-acid adsorption onto a model glass surface. This was then followed by molecular modeling of the same system in order to provide insights into the potential molecular mechanisms involved in the adsorption process. 541

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MATEMAL S AND METHOD S The adsorption enthalpy for mid-chain L-lysine adsorption onto a model silica glass surface was first measured by a Langmuir isotherm adsorption process. Poly-L-lysine hydrobromide salts with poly-L-lysine molecular weights (MW) of 1000, 1000-3000, 5000-15000, and 15000-30000 g/mol (Sigma Chemical, St. Louis, MO, USA) were used as the adsorbate and borosilicate glass microspheres with diameter between 105-125 |Lim (Cataphote Inc., Jackson, MS, USA) were used as the adsorbent. Phosphate buffered saline (pH=7.3) was used as the solvent. Poly-L-lysinesaline solutions were mixed having initial concentrations (CQ) of 0.005, 0.010, 0.030, and 0.100 mg/ml. The microbeads were methanol washed, dried, and weighed into Eppendorf vials. 1.5 ml of each concentration of poly-L-lysine/saline solution was added to each vial and rotated endover-end for 2 h in a convection oven at constant temperatures of 25, 37, 45, and 55’’C to equilibrate. Control (saline without poly-L-lysine) and experimental vials were prepared for each concentration-temperature combination in duplicate and triplicate, respectively. A ninhydrin assay (Sigma Chemical Co., St. Louis, MO, USA) was used to measure the equilibrium concentration (Ce) of poly-L-lysine in solution. The amount of poly-L-lysine adsorbed on the glass at equilibrium (qe) was determined via mass balance by comparing CQ to Ce- The adsorption isotherm data was then plotted as qe vs. Ce for each temperature and MW of poly-L-lysine, and the Langmuir isotherm equation:

QC

was best fit to the data by nonlinear regression, with Q representing the amount of poly-L-lysine adsorbed at surface saturation and a = 55.0 exp(AG/RT), where 55.0 represents the activity of water in physiologic saline, AG = change in Gibb’s free energy/mole, R = ideal gas constant, and T = absolute temperature; and where AG = AH - T AS, with AH = enthalpy change/mol and AS = entropy change/mol. The mean – 95% confidence interval of AH was directly determined for each MW of poly-L-lysine by non-linear regression using SAS (Statistical Analysis Software, SAS Institute, Cary, NC). The values of AH were obtained for each MW range of poly-L-lysine and plotted versus the degree of polymerization (DP) of poly-L-lysine (lysine residue MW = 1 2 9 g/mol), with the slope of the initial linear portion of this plot providing an estimate of the average AH value per residue of poly-L-lysine adsorbed. Following the experimental work, molecular modeling of lysine adsorption to silica glass was conducted using a CAChe Worksystem (CAChe, Beaverton, OR, USA) through a Macintosh computing environment. The geometries for silicate rings and the lysine molecule were optimized using the AMI Precise unrestricted Hartree-Fock level of molecular theory [7]. The optimized structures included the interaction of 3 to 6 member silicate rings with lysine. In these silicate rings, there are generally 2 bridging Si-O-Si bonds between adjacent tetrahedra and 2 non-bridging Si-OH bonds, or silanols. Lysine has a peptide structure [-NH-CH(R)-CO-] with the R side-group being [-(CH2)4-NH3"*"]. The lysine molecule was terminated by -H and -OH groups to satisfy its molecular structure as required for the molecular models. In the modeling process, the lysine and silica rings were brought together and the geometry of the combined structure optimized to determine its minimum energy conformation. The energy released upon lysine adsorption to the silica ring was calculated as the difference in energy between the optimized silica ring and lysine molecule taken separately compared with the energy of the optimized silica ring - lysine complex. Because the experimental portion of this study involved poly-L-lysine (and not just lysine), it was recognized that secondary structural effects of the poly-L-lysine would likely play a role in the measured adsorption process. For this reason, a second modeling study was conducted to predict the likely secondary structure of poly-L-lysine, with this data then being used to assess the

Adsorption of L-Lysine Onto Silica Glass: L.A. Latour Jr. et al.

0

20

40

60

Cp (nmol/ml)

80

100

Figure 1. Adsorption Isotherm. 1000 MW, 25*’C

0

20

40

60

80

100

Poly-L-Lysine (DP)

543

120

Figure 2. Enthalpy vs. Degree Polymerization

likely availability of the amine-terminated side-chains of poly-L-lysine to the silica surface. For this study, a 16-peptide chain of lysine (MW=2,082 g/mol) was modeled and optimized using an MM2 based molecular mechanics method [8]. RESULT S From the experimental adsorption studies, the adsorption isotherms for each poly-L-lysine MW force field were found to follow typical Langmuir behavior and were used to calculate the values of the adsorption enthalpy for each system. An example isotherm plot of qe versus Ce for 1000 g/mol poly-L-lysine adsorption onto the glass at 25 C is shown in Figure 1. The plot of AH vs. DP is shown in Figure 2. This reveals an initial linear portion, followed by an apparent plateau beginning around the DP=40 data point. The plateau evident for large DP is believed to reflect retained tertiary conformation of the poly-L-lysine molecular structure following adsorption. This effect has also been observed in polymer adsorption studies [9]. The best fit slope of the linear portion of Figure 2 is -0.23 –0.13 kcal/mol/DP (mean – 95% confidence interval), with this representing an estimate of the energy/mol released per average lysine residue adsorbed onto the silica glass surface. The molecular modeling AMI results predicted hydrogen bond formation between the sidegroup amine of lysine with the silanols of the 3-, 4-, 5-, and 6-member silicate rings with interaction energies (AH) of -4.6, +1.5,-1.1, and -6.8 kcal/mol, respectively. Figure 3 shows the results of the optimized 5-member ring-lysine structure. The relative mole fractions of 3-, 4-, 5-, and 6-member rings in silica glass have been estimated to be 0.026, 0.181, 0.275, and 0.518, respectively [10] and the Si02 concentration in the glass microspheres used in the experimental study was 72% [6]. Based upon these values, the net interaction energy of lysine with silica was calculated to be -2.6 kcal/mol. Although this value is higher than the experimentally measured adsorption enthalpy of -0.23 kcal/mol value, secondary molecular structure must also be considered. The MM2 optimization of the secondary structure of poly-L-lysine is presented in Figure 4. As shown, the analysis predicted that the molecule gently rotates with about 1 full rotation for every 8 residues. Thus, if this secondary structure is^maintained after adsorption, only about 1 or 2 out of every 8 residues may actually be accessible for interaction with the silica surface. Assuming this, the experimentally derived value may be artificially low by a factor of 4x to 8x, for a revised estimated experimental energy/mer value of -0.92 – 0.52 to -1.84 – 1.04 kcal/mole, which is reasonably close to the AMI result of -2.6 kcal/mol.

544


’ \ SILIC A X

Cten

53.5 Afi Silicon

P ^’2.511 A C t

Figure 1. Ternary compositiona l diagram where the composition s BV-01, BV-11 and BV-15 have been are shown. The glass transition temperature s obtained by dilatometri c methods can be seen in Table II . Such temperature s have been used to give the annealing treatmen t to the glasses. The results of the microhardnes s indentatio n tests carried out are summarized in Tables III and IV . It can be noticed that the Vickers hardness for these materials ranges betwee n 2.74 and 3.57 GPa and their Young’s moduli betwee n 35 and 93 GPa. By means of a Vickers indentatio n and by measuring the radial cracks produced, the fracture toughness was also evaluate d and it ranged betwee n 0.79 and 3.45 MP a m’’’^. In order to obtain a bette r understanding , the critical strain energy release rate was also evaluate d in plain strain as:

G.=%^(1-.’)

(3)

where v is the Poisson ratio which for most glasses is approximatel y 0.2. The values obtained are also shown in Table IV . Table II . Glass transition temperature s of the glasses studied.

1

1

Glass

IîlQ

1L

BV-01 329. 1

BV-11 445.1

_

BV-15 367.1

Table III . Vickers Hardness (HV) and Young’s modulus (E) for the different glasses studied.

1 1

Glass BV-01 BV-11 BV-15

HVlGPa l 2.74 – 0.07 3.57 –0.08 3.06–0.0 5

E (GPa) 1 35.62 – 2.35 92.82 – 5.26 53.04 + 7.04 1

1 1

556


Table IV. Fracture toughness of the different glasses studied. 1

Glass BV-01 BV-11

L BV-15

Kic (MPa m^^^) 0.8210.05 0.79 –0.03 3.45 –0.12

Gic(kJ.m-’)

1

1.81x10-^ 0.65 X 10-^

21.54x10’^

1

The present results can be interpreted in terms of the different CaO and Na20 contents and their bonding to the P2O5 network. In fact the structure of vitreous phosphoric oxide is described by chains as in the case of a polymer. Both Na and Ca ions have the ability to break the P2O5 chains, but whilst Na does not produce any strong bonding between lateral chains, the effect of Ca is to link them [8]. In other words, whilst Na breaks and weaken the P2O5 chains, Ca links the chains together. This interpretation gives ground to understand that with increasing Ca content the glasses will become stiffer, harder and with a higher glass transition temperature. In fact this is in complete agreement with the results obtained. It can be noticed that as the CaO content increases: BV-01, BV-15 and BV-11, the glass transition temperature, the Vickers hardness and the Young’s modulus increase. However, fracture toughness seem to correlate with the P2O5 content. In fact, it seems also that Ca may act as an embrittlement agent, since for a minor variation in P2O5 for BV01 and BV-11, and a large increase in Ca content, the strain energy release rate decreases by a factor of 3. However, further research should be required in order to provide a proper explanation, since fracture toughness is also very sensitive to the defect population in the glass. ACKNOWLEDGEMENT S The authors of the present study are grateful to the CICYT for providing financial support through the project MAT96-0981. REFERENCE S l.C.F. Drake and M. Graham, Inorganic glasses as slow release herbicides and fungicides, Chemical Society, Burlington House, London, (1976). 2. J.Bumie, T. Gilchrist, S.R.L Duff, C.F. Drake, N.G.L. Harding and A.J. Malcolm, "Controlled release glasses for biomedical uses", Biomaterials, 2 (4), (1981), 244-245. 3. J.Bumie and T.Gilchrist, "Controlled release glasses: A new biomaterial", in Ceramics in Surgery , Ed. P.Vicenzini, Elsevier, The Netherlands, (1983), 169-176. 4. A. Afonso, J.D. Santos, M. Vasconcelos, R. Branco and J. Cavalheiro, "Granules of osteoapatite and glass-reinforced hydroxyapatite implanted in rabbit tibiae", J. Mater. Science: Materials in Medicine, 7, (1996), 507-510. 5. G.R.Anstis, P.Chantikul, B.R.Lawn and D.B.Marshall, "A critical evaluation of indentation techniques for measuring toughness: I", J. Amer. Ceramic Soc, 64 (9), (1981), 533-538. 6. D.B. Marshall, T. Noma and A.G. Evans, "A simple method for determining elastic-modulus-tohardness ratios using Knoop indentations measurements", Communications of the Amer. Ceramic Soc, (1982), C-175-176. 7. M.Ontan6n,F.J.Gil,A.Casinos,F.Guiu and J.A.Planell, "Young’s modulus and fracture toughness of cortical bone evaluated by means of indentation techniques", in Biomaterial-Tissue Interfaces, Eds.: P.Doherty et al., Elsevier, Amsterdam (1992), 171-177. 8. J.E. Shelby, W.C. Lacourse and A.G. Clare, "Engineering Properties of Oxide Glasses and Other Inorganic Glasses", in Ceramics and Glasses, Engineered Materials Handbook, Vol. 4, ASM, USA, (1991), 845-857.


SILICO N IN CONNECTIV E TISSUE : SEMI-EMPIRICA MODEL S

L MOLECULA R ORBITA L

Keith D. Lobel\ Jon K. West\ and Larry L. Hench’’^ Department of Materials Science and Engineering, University of Florida, Gainesville, FL 32611 ^Department of Materials, Imperial College, University of London, Prince Consort Rd.

ABSTRAC T Semi-empirical molecular orbital modelling is used to predict the chemistry of interactions between soluble silicon, Si(0H)4, and biological macromolecules of connective tissues which are known to possess high concentrations of silicon. Results show a particularly favorable low-energy reaction pathway which forms ester-like linkages between silanol groups and carbonyl groups, with a reaction barrier of only 0.2 kcal/mol and a systemic stabilization of 4.6 kcal/mol. The effect of organic templates on the nucleation and growth of hydrated silica is also demonstrated. The potential impact of these data on the 1) osteogenic properties of silicate-based bioactive glasses and 2) employment of silica-rich mineral exoskeletons by marine organisms is discussed. KEYWORD S

silicon, silica, molecular models, protein, adsorption

INTRODUCTIO N The structural role of hard connective tissues in higher order animals is met by the presence of rigid carbonated hydroxyapatite crystals embedded within a flexible matrix of type I collagen and ground substance (interwoven proteoglycans and noncollagenous glycoproteins). However, both hard and soft connective tissues have been shown since the work of Carlisle [1], and Schwarz and Milne [2] to contain relatively high quantities of silicon (Si), which is believed to play an important structural role in these tissues. Biological evidence suggests that elemental silicon exists as cross-linking chemical species between chains of glycosaminoglycans, the major structural polysaccharides of vertebrates [3]. Little is known, however, of the chemical binding sites, or chemical/metabolic pathways, which result in linking the organic and inorganic constituents. The prevalence of hydroxy 1 and carboxyl groups of both proteins and saccharides in these tissues make such sites potential candidates for reaction with elemental silicon [1,2,4], and warrantfiirtherinvestigation. Certain types of lower order plants, such as diatoms and radiolarians, clearly demonstrate the use of polymerized hydrated silica for growth of inorganic structural exoskeletons, which are also known to be intimately associated with an organic matrix [5,6]. Although the Si is present here as bulk-form silica, the interactions between organic and inorganic phases may be similar to those in connective tissues of higher animals. Other physiological roles of silicon are implicated by the osteogenic characteristics of silica-based zeolites, zeolite extracts, sodium silicate solutions and certain compositional ranges of bioactive glasses [7]. The mnate complexities of these biological systems makes it difficult to identify the salient modes of interaction between silicon and biomolecules of the host. While its limitations must be recognized, molecular orbital modelling offers an additional investigative tool for studying the basic science of biological interfaces which is able to circumvent the inherent problems of real-system analysis. 557

558


METHOD S Several types of dehydrolysis reactions were simulated using the semi-empirical AMI computational model. The AMI method is a self consistent field method of calculation which is parameterized for molecules containing silicon, carbon, oxygen, nitrogen, and hydrogen. A recent review justifies the application of this model to the systems at hand [7]. In an extension of earlier MO modelling work by the authors [8,9] reactions between silanol groups and various functional groups of amino acids and sugars are investigated. Reaction energy pathways were calculated and transition states for the molecular mteractions were determined and verified with Unrestricted Hartree Fock (UHF)-level force calculations. Care was taken to insure that terminal ends cf opposite R-groups did not meet each other and contribute a false component of systemic stabilization. Silicon cross-links were modelled by reacting two similar organic molecules, for example serine (Ser), with two silanol groups of a silicic acid molecule: 2 Ser + Si(0H)4 > Ser-0-Si-O-Ser + 2H2O In addition, the energetics of organic template stabilization on the process of silicification was studied by modelling the sequential condensation of four silicic acid molecules to form a tetrasiloxane ring covalently bound to the template, as described elsewhere [9]. In this study, efforts were concentrated on hydroxyl and carboxylic acid groups of both sugars and amino acid side-chains, since these have been proposed as the reaction sites of the organic host material, in both connective tissue (bone and cartilage) and organic templates of siliceous frustules [1-4]. Figure 1 demonstrates the possible linking sites of Si(0H)4 between two hyaluronic acid monomers. Protein template models were constructed to mimic the p-sheet structure, while saccharide templates were given free mobility. (The distance between two cw-silanol groups of a tetrasiloxane ring is 4.68 A, which is within 0.03 A of the interpeptide distance of the anti-parallel p-sheet). The structural stability of a protein p-sheet results from a combination of regular interpeptide hydrogen bonds and large barrier to peptide bond rotation [3]. These conformational effects are not as prominent in the polysaccharides, due in part to pure hydrocarbon backbones. RESULT S AND DISCUSSIO N The process of silonate bridge formation and biosilicification were modelled through, respectively, two and six 4-step dehydrolysis reactions. These consistent stages of each reaction include 1) hydrogen-bonding of reactants, 2) saddle point (activation state), 3) low-energy pentacoordinate metastable transition state, and 4) products. Similar mechanisms were found in other studies [10,11]. Complexation of silanols with hydroxyl groups form ether-like bonds (i.e. C-O-Si) while carboxyl groups formed ester-like bonds (i.e. C-OO-Si) (see Figure 1). Figure 2 shows a bar graph of the activation barriers and net degrees of stabilization fir each of the silonate bridge models. Only the systems that formed ester-like bonds resulted in products that were more thermodynamically stable than their reactants. The dramatic difiFerence between barriers of silanol complexation with glutamic acid (17.8 kcal/mol) and an acidic monosaccharide (0.2 kcal/mol) are believed to be due to different mechanisms of reaction rather than the different substrate structures. Figure 3 shows saddle points of the two reaction complexes. During ligation of silicic acid to glutamic acid, the resonance form shown above, Figure 3 a, was maintamed, and the -OH oxygen of the carboxyl group ligated with silicic acid. During reaction with the acidic saccharide, however, the carboxyl group adopted the resonance form shown in Figure 3b, and the non-protonated oxygen ligated with silicic acid. The more favorable energetics of the system in 3b may be due, in part, to 1) a more readily donated hydrogen from the cationic oxygen of 3b, and 2) the fact that simultaneous donation of H^ in 3 a to form water, and ligation with Si via the same -OH is not as sterically favorable as the latter case, where these two events occur via separate oxygens. Figure 2 also shows the activation barriers and net degrees of stabilization for five models of template-mediated biosilicification. Formation of a tetrasiloxane ring in the absence of an

Silicon in ConnectiveTissue: Semi-Empirical Molecular Orbital Models: K. Lobel et al. 559

25x Biosilicificatio n

NHCOCH3

C H , OH

«^ ,-# 1

C H j OH

^

0

’

Si(OH) 1

Models

1

COO"

1 ^

Sllonate-Bridg e ^

0f

2

Si(OH) 1

.^^

2

1

COO"

Barrier

NHCOCH3 Figure 1. Two hyaluronic acid monomers with two potential silonate cross-links

Stabilization

Figure 2. Thermodynamics of silonate bridge formation and biosilicificatio n models.

O-H

R-c 80 %

t

"^Wate r Formatio n

Resonanc e O-H

R-C %.

20 % -Wate r Formatio n Figure 3. Saddle structures for silonate bridge formation betwee n acidic monomers, (a) complexatio n to glutamic acid through the more prevalent (80%) resonance form of COOH . (b) complexatio n to an acidic monosaccharid e through the less prevalent (20%) resonance form of COOH .

560


organic template has the greatest thermodynamic stabilization (-41.8 kcal/mol). Regulation by acidic organic templates was found to be more favorable than hydroxyl-rich templates in terms of both net stabilization and activation barriers. This data supports the theory of organic matrix mediated biosilicification: the initial nucleation of a silica-rich phase may occur via acidic sites on an organic template, as suggested by the low activation barrier for this system. Subsequent autopolymerization of silicic acid may proceed without further matrix influence to build the siliceous skeleton by the matrix-free model. Such a model meets the necessary role of lowering the energy of nucleation for biomineralization [5,12]. A similar functional role of collagen and sulfonated macromolecules of ground substance during the mineralization of bone is believed to exist [13]. Unfortunately the AMI model is currently not parameterized for calcium, making theoretical analysis difficult. Advances in molecular orbital chemistry will likely make fiulher analysis possible in the near future. A recently developed theory to account for the enhanced osteogenic properties of Class A bioactive glasses and glass-ceramics relative to Class B materials (e.g. hydroxyapatite) focuses on the differences in soluble silicon production [14]. The biomolecular and/or cellular mechanisms by which silicon may promote tissue production remams to be identified - indeed, a potential structural/metabolic dichotomy should be recognized. Whether the element induces rapid proliferation by merely supplying one of the necessary biulding blocks of connective tissue [1,2], or by actively triggering an increased metabolic activity of formative cells (e.g. chondroblasts, osteoblasts) [15], one can assume that highly specific interactions with biomolecules of these tissues will ultimately dictate the biological response. A more complete understanding of these interactions may help reveal the key to the unique behavior of silicate-based bioactive glasses. ACKNOWLEDGMEN T The authors would like to thank the Air Force OflBce of Scientific Research for financial support of this work, under Grant No. F49620-92-J-0351. REFERENCE S 1. Carlisle, E.M., Science 1972, 178, 619-21. 2. Schwarz, K. and Mihie, D.B., Nature1972, 239, 333-4. 3. Mathews, C.K. and van Holde, K.E., In: Biochemistry,Benjamin/Cummings Pub. Co., Redwood City, CA. 1990, 287. 4. Hecky, R., Mopper, K., Kilham, P. & Degens, E., Mar. Biol 1973, 19, 323-331. 5. Perry, C.C, In: Biomineralization:Chemicaland BiochemicalPerspectives,VCH, London 1989, 223-56. 6. Volcani, B.E., In: Silicon and SiliceousStructuresin BiologicalSystems,SpringerVerlag,New York 1981, 157-200. 7. Hench, L.L. and West, J.K., Life ChemistryReports1996, 13, 187-241. 8. Hench, L.L. and West, J.K., In: Bioceramics6,Buttenvorth-Heinemann, Oxford, 1993, 35-40. 9. Lobel, K.D., West, J.K., and Hench, L.L., Mar. Biol. 1996, 126, 353-60. 10. Burggraf, L.W., Davis, L.P. and Gordon, M.S. In: UltrastructureProcessing of AdvancedMaterials,Wiley, New York, 1992, 47-55 11. West, J.K. and Wallace, S., J. Non-Cryst.Solids1993, 152, 109-117. 12. Mann, S. et al.. Science1993, 261, 1286-92. 13. Lowenstam, H.A. and Werner, S., In: On Biomineralization, Oxford Univ. Press, New York 1989. 14. Hench, L.L., In: Bioceramics7, Buttenvorth-Heinemann, Oxford, 1994, pp. 3-14. 15. Keeting, P.E. et al., J. ofBone and Min. Res.1992, 7, 1281.


TH E EFFEC T OF HEA T TREATMEN T ON BON E BONDIN G ABILIT Y OF ALKALI-TREATE D TITANIU M Shi gem Nishiguchi \ Takashi Nakamura\ Masahiko Kobayashi \ Wei-Qi Yan\ Hyun-Min Kim^, Fumiaki Miyaji ^, and Tadashi Kokubo^ ^Department of Orthopedic Surgery, Faculty of Medicine, Kyoto University, Shogoin-kawaharacho 54, Sakyo-ku, Kyoto 606-01, Japan ^Department of Material Chemistry, Faculty of Engineering, Kyoto University, Yoshida-honmachi, Sakyo-ku, Kyoto 606-01, Japan ABSTRAC T The purpose of this study is to evaluate the bone-bonding ability of alkali-treated titanium without heating. We made three types of titanium plate. 1) control group; pure titanium 2) alkali-treated group 3) alkali-and heat-treated group. The reaangular plates were inserted transcortically into the proximal met^^hyses of both rabbit tibiae. The tensile failure loads between implant and bone were measured by a detaching test. The tensile failure loads of alkali-and heat-treated group was 2.67 kgf and 4.13kgf, at 8 and 16 weeks, respectively. In contrast, those of control group and alkali-treated group were nearly 0 kgf even at 16 weeks. Histological examination revealed that alkali-and heattreated group was in direct contact with bone, but the other groups have thin intervening fibrous tissue between the implants and bone. A previous in vitrostudy demonstrated that alkali treatment without heating provide titanium an ability to form apatite on its surface. However our data showed alkali-treated titanium had no bone bonding ability. This discrepancy was due to the unstable surface reactive layer of alkali-treated titanium, which might be lost during preservation or implantation. In conclusion, both alkali and heat treatment are essential for preparation of bioactive titanium in practical use. KEYWORD S titanium, alkali treatment, heat treatment, bioactivity , bone-bonding IIVTRODUCTIO N In vitrostudy Kokubo et al showed that via a simple chemical treatment of alkali treatment and heat treatment, titanium and its alloys formed bonelike âtite on their surface in simulated body fluid (SBF), which has ion concentration neariy equal to human blood plasma(1-4) In those articles, alkali-treatment alone also can provide titanium an ability to form apatite. Apatite formation on the material surface is believed to be a prerequisite for its bioactivity, that is, direct bone bonding.(5) In vivo study, Yanet al reported that alkali-and heat-treated titanium can bond to bone directly , and also showed that titanium treated alkali plus heat and SBF soaking has bone-bonding ability.(6,7) If alkali-treated titanium bond to bone without heat treatment, preparation of bioactive titanium will become much easier. And the effect of heat treatment to alkali-treated titanium on bone bonding ability is unknown. The purpose of this study is to investigate whether alkali-treatment alone can provide titanium bioactivity, and to investigate the influence of heat treatment on bone bonding strength. MATERIAL S AN D METHOD S Implant preparation We made three kinds of rectangular titanium plates (size; 15X10X2 mm). 1) control group; Commercially pure titanium plates (Kobe Steel Co., Kobe, Japan) were abraded 561

562


with #400 diamond paste and washed with distilled water. And they were dried at room temperature. 2) alkali-treated group ; Pure titanium plates were soaked in 5 M NaOH aqueous solution for 24 hours. After the alkali treatment, the plates were gently washed with acetone and distilled water, and dried at 40*C for 24 h in an air atmosphere. 3) alkali-and heat-treated group ; After alkali-treatment, the plates were heated to 600*t at a rate of 5 C/min. in an elearic furnace, kept at a given temperature for Ih, and cooled to room temperature in the furnace. Implantation The implants were conventionally sterilized with ethylene oxide gas. They were implanted into the metaphyses of the tibiae of mature male Japanese white rabbits. The surgical methods was the same as that reported previously.(8) Using a dental burr, a 16X2 mm hole was made from the medial cortex to the lateral one parallel to the longitudinal axis of the tibial metaphysis. After irrigation of the hole with saline, the titanium plates were implanted in the frontal direction, perforating the tibia and protruding from the medial to lateral cortex. Four rabbits of each group were sacrificed at 8 and 16 weeks after the operation. In this study the guideline for animal experiments of Kyoto University was observed. Measurement of the detaching failure load After sacrifice, segments of proximal tibial metaphyses containing the implanted plates were cut out and prepared for the detaching test.(Figure 1) Traction through hooks holding the bone segments was applied vertically to the implant surfaces at a crosshead speed of 35 mm/min using an Instron-type autograph (Model 1011, Aikoh Engineering Co., Ltd., Nagoya, Japan). (8) The detaching failure load was measured when the plate was detached from the bone. If the plate was detached before the test, the failure load was defined as 0 Kgf. Data were expressed as mean – standard deviation (SD) and assessed using a one-way ANOVA. Differences at p < 0.05 were considered to be statistically significant. Histological Examination After the detaching test, specimens were fixed in 10% phosphatebuffered formalin and dehydrated in serial concentratbns of ethanol. Then, they were embedded in polyester resin. Sections 500 u m thick were cut with a band saw (BS-3000, EXACT cutting system, Norderstedt, Germany) perpendicular to the axis of the tibia, and were ground to a thickness of 150-180 u m for CMR and Giemsa surface staining using a grinding-sliding machine (Microgrinding MG-4000, EXAKT). Several 500 u m sections were polished with diamond paper and coated with a thin layer of carbon for observation by a scanning electron microscope (S-800, Hitachi Co. Ltd., Tokyo, Japan) attached to an energy-dispersive X-ray microanalyzer (EMAX-3000, Horiba Co. Ltd., Kyoto, Japan) (SEM-EPMA). RESUL T Detaching test In detaching test breakage always occurred at the bone-plate interface. The detaching failure load of each material at 8 and 16 weeks after implantation are summarized in Table 1. At 8 weeks after the operation the failure load of alkali-and heat-treated group was significantly higher than those of the other groups. At 16 weeks after the implantation, alkali-and heat-treated group showed a failure load of 4.13 kgf This value was significantly higher than that of the same group at 8 weeks and those of alkali- treated group and control group at the same weeks. The failure loads of alkali-treated and control group at 16 weeks did not differ significantly from those at 8 weeks. Table 1

Detaching test

failure loads (kgf; mean – SD) 8w 2.71+1.47 alkali-and heat-treated group (n=8) 0.52+0.52 alkali-treated group(n=8) control group(n=8) 0.02+0.03

16w 4.13 –1.25 0.49+0.38 0.33+0.36

Effect of Heat Treatmenton Bone Bonding Ability of Alkali-Treated Titanium:S. Nishiguchi et al.

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Histological examination Histlogical examination by Giemsa surface staining and CMR revealed that alkali-and heat-treated group showed direct contact between bone and plate without any intervening fibrous tissue in all samples at 8 weeks(Figure 2). And at 16 weeks amount of bone directly contacted to the plate increased. In contrast, there wasfibroustissue layer between bone and the implant in both control and alkali-treated groups. At 16 weeks new bone formation increased around the plate, but intervening layer still remained in these groups SEM examination showed almost same findings as Giemsa and CMR Control and alkali-treated groups had a gap between bone and titanium, while alkali-and heat-treated plate contacted bone directly. Alkali-and heattreated group showed thin new bone layer formed on intramedullary portion of titanium far from cortex even at 8 weeks. Thisfindingwas not seen in alkali-treated and control groups. DISCUSSIO N Alkali- and heat-treated titanium was reported to bond to bone and supposed to be clinically applicable as orthopaedic implant material. (6.7) If alkali-treatment without heat treatment could induce bioactivity on titanium ,it would become possible to prepare bioactive titanium implant with soaking titanium implant in NaOH solution even at the operation theater. In this study, the bonding strength of alkali-and heat-treated titani um increased to 4.17kgf at 16 weeks by a detaching test. In contrast, alkali-treatedfitaniumshowed almost no bonding. And pure titanium did not bond to bone even at 16 weeks in accordance with the previous report. (9) Kim et al reported that alkalitreated titanium without heating had an ability to form apatite on its surface in SBF just like alkali-and heat-treated one.(3) And they also showed that once apatite layer formed, the tensile strength of the apatite layer and alkali-treated titanium is comparable to that of alkali-and heat treated one. (10) Thus if apatite layer forms, there will be no difference between bonding strength of alkali-and heat treated titanium and that of alkali-treated one. In this study, however, the bonding strength of alkali-treated titanium was significantly lower than that of alkali-and heat-treated one. This discrepancy can be explained as follows. Although alkali titanate hydrogel layer forms on titanium when soaked in alkali solution, this layer is so unstable as to lose its apatite forming ability during preservation or implantation. Throughout this study the implants were treated manually like as Figure 1. Detaching test

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orthpaedic implants and not treated so roughly. Although alkali-treated titanium has an ability to form apatite in SBF, it does not serve as bone bonding material in practical use. If alkali-treated titanium bond to bone, preparation of bioactive titanium would become easier. However this is not recommended from these results. In practical use both alkali and heat treatment is essential for inducing bioactivity on titanium. REFERENCE S 1. Kim, H. M., Miyaji, R, Kokubo, T., Nakamura ,T. J.Biomed.Mater.Res., 1996, 32,409-417. 2. Kokubo, T., Miyaji, F., Kim, H. M., Nakamura, T. JAm.Ceram.Soc.1996,79, 1127-1129. 3. Kim, H. M. , Miyaji, P., Kokubo, T., Nakamura, T. 7. Ceram.Soc. Japan, 1997, 105, 111-116. 4. Kokubo, T., Kushitani, H., Sakka, S., Kitsugi, T., Yamamuro, T. J.Biomed.Mater.Res.1990,24, 721-34. 5. Kokubo, T., Ito, S., Huang, Z. T., et al. J. Biomed.Mater.Res. 1990,24, 331-43. 6. Yan, W. Q., Nakamura, T., Kobayashi, M., Kim, H. M., Miyaji, P., and Kokubo, T. J.Biomed. Mater.Res., to appear 7. Yan, W. Q., Nakamura, T., Kobayashi, M., Kokubo, T., Kim, H. M., Miyaji, P. In: Bioceramics volume 9, Elsevier, Oxford, 1996, 305-308. 8. Nakamura, T., Yamarumo, T., Higashi, S., Kokubo, T., Itoo, S. J.Biomed.Mater.Res.1985, 23, 631-648. 9. Takatsuka, K., Yamamuro, T., Nakamura, T., Kokubo, T. J Biomed.Mater.Res. 1995,29, 15763. 10. Kim, H. M., Miyaji, P., Kokubo, T., Nakamura, T. In: Bioceramicsvolume 9, Elsevier, Oxford, 1996, 301-304. Pigure 2. Giemsa surface staining image at 8 weeks; a.Alkali-and heat-treated group. Titanium implant has direct contaa with bone. b. Alkali-treated group. Thin intervening layer exists at the interface between bone and titanium. (X200)

B=bone T=titanium I=intervening layer


THERMA L PROCESSIN G OF COMPAC T BOVIN E BON E G. Vargas, M. Mendez, J. Mendez, and J. Lopez. CINVESTAV-IPN Unidad Saltillo, Apartado postal 663, 25000 Saltillo, Coahuila, Mexico.

ABSTRAC T In the present work, chemical and structural transformations undergone by compact bovine bone thermally processed up to 1500 C were studied using x-ray diffraction, scanning electron microscopy, and optical microscopy. The results showed that most of the organic matter was burned at temperatures below 500 C. The residual carbon content of bone was ~1 wt% at this temperature. The carbonate hydroxyapatite phase remained stable up to 600 C. At 1000 C, hydroxyapatite plus 0.6 wt% Na, 0.7 wt% Mg, and -0.1 wt% C, were found. Up to this temperature, 37 % of the initial weight of the boiled and skinned compact bovine bone had been lost. Between 1200-1500 C, mixtures of a-tricalcium phosphate, tetracalcium phosphate and hydroxyapatite, were observed. At these temperatures, the relative amount of each phase was affected by heating time and cooling rate. KEYWORDS : Bone calcination, calcium phosphates, hydroxyapatite, carbonate hydroxyapatite. INTRODUCTIO N When bones are thermally processed, water, living cells and other organic components, are removed, and inorganic matter suffers chemical and structural changes. Bovine bone calcining is a promissory process to obtain different calcium phosphates for biomedical applications such as drug delivery systems [1], bone repair and reconstructive surgery [2], as well as for manufacturmg of cements used in temporary or permanent dental and orthopaedic implants fixation [3]. Several research works have been carried out on bone calcmmg [4,5,6]. However, as different types of calcium phosphates may show different bioactivities inside a living being, it is necessary to fully understand their formation process during bone calcining. Thus, further studies about the chemical and structural transformations taking place during bone thermal processing, are required. MATERIAL S AND METHOD S Thin round sections of compact bovine leg bones were used. The first stage of bone processing was degreasing in boiling water. Subsequently, the bones were skinned before the remaining organic matter was burned at 300 C. After burning, each bone section was thermally processed for up to 24 hours at temperatures between 500 and 1500 C. Once cooled down to room temperature, the calcined bones were separated into two parts, one of which was subjected to a milling process, until the particles were smaller than 5|im, to carry out X-Ray Diffraction studies. The other bone piece was prepared using standard ceramographic techniques in order to carry out 565

566


Studies of optical and scanning electron microscopy. More detailed information about the experimental equipment and techniques employed can be found in a previous work [7]. RESULTS AND DISCUSSION The following results were obtained from a previous work involving the chemical and microstructural characterisation of compact bovine bone processed up to 1000 C [7]: 1. Bone is dehydrated, fat is burned, and volatile matter is eliminated below 350 C, involving a total loss of-12.5% of the initial weight of the compact bovine bone. 2. Thermal cycles at 500 C yield products with up to ~lwt% of residual carbon, involving a total weight loss of ~34wt%. 3. The thermal treatment should be carried out between 500 and 600 C if carbonate hydroxyapatite has to be preserved, because the occurrence of carbonate decomposition was observed to take place between 600 and 700T. 4. The residual carbon content at SOOT is -0.14 wt%. 5. By subjecting powdered samples to x-ray diffraction, as well as through the examination of monolithic samples employing optical and scanning electron microscopy, only a single homogeneous phase was detected in bone processed at temperatures between 700 and 1000 C, which was identified as hydroxyapatite. 6. After treatment at lOOOT, 0.6 wt% Na, 0.7 wt% Mg, and 0.12 wt% C, were found. The weight of bone calcined at this temperature was approximately 63% of the initial weight of the boiled and skinned compact bovine bone.

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Thermal Processing of Compact Bovine Bone: G. Vargas et al.

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Figure 2. XRD pattern (a) and Optical photomicrograph (b) of bone calcined at 1500 C for 2 hrs. In the present work, the results obtained from further studies carried out at temperatures between 1200 and 1500 C, are reported. According to the x-ray diffraction pattern shown in Figure 1(a), hydroxyapatite was the main phase remaining after bone processing for up to 24 hours at 1200 C. However, as it is shown in Figure 1(b), at this temperature the incipient formation of a new phase was noticed on the periphery of the Haversian channels under the optical and the scanning electron microscopes. Later on, EDS microanalysis on the SEM allowed to identify the new phase as tricalcium phosphate. The rate of formation of this phase was very slow at 1200 C. Thus, the presence of tricalcium phosphate could not be detected neither by x-ray diffraction, even after 24 hours of treatment, nor by differential thermal analysis [7], although it was followed, employing monolithic bone samples, by sequential observations made under the optical microscope. After 2 hours at 1500 C, and allowing the samples to cool down rapidly outside the furnace, the presence of three different phases (hydroyapatite, tricalcium phosphate and tetracalcium phosphate) was confirmed by x-ray diffraction, as shown in Figure 2(a). The EDS spectrum obtained on the SEM for the light grey phase observed in Figure 2(b) is given in Figure 3(a). The semiquantitative determination of the Ca/P molar ratio obtained from this EDS spectrum gave a value of 1.6, which is close to the stoichiometric value of 1.667 corresponding to hydroxyapatite. The dark grey phase observed in Figure 2(b) showed a Ca/P molar ratio of 1.3 from the EDS pattern displayed in Figure 3(b). Since a suitable standard to carry out a more accurate determination could not be employed, this value was assumed to be reasonably close to the stoichiometric value of 1.5 corresponding to tricalcium phosphate. A small amount of Na was also detected in this phase. From the XRD diffraction pattern shown in Figure 2(a), and from the EDS studies, the white phase observed in Figure 2(b) was inferred to be tettacalcium phosphate. For treatment times longer than 24 hours at 1500 C, only mixtures of tricalcium phosphate and tetracalcium phosphate were found. On the other hand, by using slow cooling rates, the reversion into oxyapatite of some of the tticalcium phosphate and tetracalcium phosphate formed was observed, which was in agreement with previous observations made by Gottschling et al [8].

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REFERENCES 1. Itokazu M., Matsunaga T., Kumazawa S. and Wenyi Yang, Journal of Applied Biomaterials, 1995,6, 167-169. 2. Ricci J.L., Blumenthal N.C., Spivak J.M. and Alexander H., J. Oral Maxillofac. Surg., 1992, 50, 969-978. 3. OonishiH., Biomaterials, 1991, 12, 171-178. 4. Webster A. V., Cooper J. J., Hampson C. J. and Cubbon R. C. P., British Ceramic Transactions, 1987,86,91-98. 5. Hill R. G., Webster K., May C. and Mandal A., British Ceramic Transactions, 1994, 93, 16-20. 6. Cooper J. J., British Ceramic Transactions, 1995, 94, 165-168. 7. Vargas G., Mendez M., Mendez J. and Lopez J., sent to Journal of Applied Biomaterials. 8. Gottschling S., Kohl R., Engel A., Oel H. J., Bioceramics: Materials and Applications, Ceramic Transactions, 1995, 48, 201-213.

EVALUATIONS METHODS AND NEW APPLICATIONS



CHARACTERIZATIO N OF SYNTHETI C AND BIOLOGICA L CALCIU M PHOSPHAT E MATEIUAL S BY MICRO-RAMA N SPECTROMETR Y G. Penel^ ^ G. Leroy\ G. Couraot^ andE.Bres^ Û 279 INSERM 1 rue du Prof. Calmette, 59019 Lille Cedex, France; ÛFR de Chirurgie-Dentaire, place de Verdun 59000 Lille, France; ^ CNRS (URA583)-INRA-LNSA, 78350 Jouy en Josas; \SPES URA CNRS 234 USTL Bat.C6 59655 Villeneuve d’Ascq. France ABSTRAC T In this study the composition of biological calcium phosphates obtained in vivo (enamel, dentine and bone, implants of bone substitutes) or in vitro (resorption cavities) was investigated by micro-Raman spectrometry and compared to synthetic compounds. This technique provides high resolution molecular informationns in a non-destructive manner and without possible artifacts due to specimen preparation. KEYWORD S MicroRaman spectrometry - calcium phosphate - biomaterials - cell culture - prothesis. INTRODUTIO N Ciystallographic and chemicals investigations of calcium phosphate of biological interest are extensively performed with different methods (chemical analysis. X-ray diffraction, infrared spectrometry, MAS-NMR, TEM, etc.; [1-5]) that allowed to know their composition and structure, but numerous unresolved problems persist about the fine characterisation of the apatites. It has been shown that Raman spectroscopy, mainly focused to the vibration modes of P04^", OH’, HP04^’ and COa^’, can provide complementary informations in a non-destructive manner. In the biological calcium phosphates however this analysis is difficult because of the fluorescence caused by the organic components. This fluorescence can be reduced using high spatial resolution micro-Raman spectrometry. The purpose of the present study was to investigate by micro-Raman spectrometry the composition of biological calcium phosphates obtained in vivo or in vitroas compared to synthetic apatites. METHOD S Micro-Raman analysis were all performed with a OMARS 89 and a LABRAM microspectrometers from DELOR (Lille-France). RESULT S Synthetic bioapatites. Apatites of biological interest are mainly represented by tricalciumphosphate p (p-TCP), hydroxyapatite (GHAp) and carbonated apatite (CarAp). Micro-Raman spectra of these compounds are shown in figure 1. The spectra are dominated 571

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(d) Figure 1. MicroRaman spectra of p-TCP (a), hydroxyapatit e (b) and AB type (c) A type (d) carbonated apatites. by the modification s of the very strong band (964 cm’^) of hydroxyapatite . The carbonate substitution or the complete lack of hydroxyl ion give raised to a frequenc y shift and split of this band, respectively . Additionally, the A type and B type carbonate substitutio n are separatly identifie d by two bands at 1107 and 1070 cm’^ respectively . Biological apatites Enamel, dentine and bone. Micro-Raman spectra of enamel, dentine and bone are shown in Fig. 2. As compared to the spectra of synthetic apatites the mineral part of dentine and bone are very similar, whereas enamel exhibit a very well-crystallize d structure and a low carbonate substitutio n and acidic phosphate content [6-10]. Hydroxyapatite and p-TCP implants in rabbit bone. Micro-Raman spectra of the implants are shown in Fig. 3. No changes were observed , as compared to the spectra of the synthetic apatites (Fig. 1) before implantation . No intermediat e phase is observed betwee n bone and implant material. (a)

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Characterizationof Calcium PhosphateMaterials by Micro-Raman Spectometry:G. Penel et al.

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SUMMAR Y This study shows that micro-Raman spectrometry allows to strictly identified the different apatites contained in biological samples, by comparaison with the spectra obtained in synthetic compomids. Fmther studies will be required to characterize organic components. REFERENCE S 1. Arends J., Christoffersen J., ChristoflFersen M.R., Eckert H., Fowler B.O., Heughebaert .C, NancoUas G.H., Yesinowski J.P. and Zawacki SJ. J. of CrystalGrowth 1987,84,515-532. 2. Kleebe H.J., Bres E., Bemache-Assolant D., and Ziegler G. J.Am. Ceram. Soc,1997, 80, 37-44. 3. Elliott J.C. In Structureand chemistryoftheapatitesand othercalcium orthophosphates. Studiesin inorganicchemistry,Elsevier, London 1994, 191-301. 4. Penel G., Leroy G., Key C, Sombret B., Huvenne J.P. and Bres E. J. Mater.Sci. : Mater, in Med.,1997 {in press). 5. Bertoluzza A., Cacciari S., Tinti S., Vasina M. and Morelli M.A. J. Mater.Sci. 1995, 6, 76-79. 6. Fateley W.G., McDevitt N.T. and Bentley F. Appliedspectroscopy1971, 25, 155-173. 7. Tsuda H. and Arends J.J. Dent. Res.1994,11, 1703-1710. 8. Nelson D.G.A. and Williamson B.E. Aust.J. Chem. 1982, 35, 715-727. 9. Rehman I., Smith R., Hench L.L. and Bonfield W. J.ofBiomed.Mater.Res. 1995, 29, 1287-1294. 10. Sauer G.R., Zunic W.B., Durig J.R. and Wuthier R.E. Calcif Tissue.Int.1994, 54, 414420.


BIOLOGICA L EVALUATIO N OF GLAS S REINFORCE D HYDROXYAPATIT BY FLO W CYTOMETR Y

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M.A. Lopes^’2, J.C. Knowles 3, K.A. Hing^ , J.D. Santos^, FJ. Monteiro^ and I. Olsen’* ÎRC in Biomedical Materials, Queen Mary and Westfield College, Mile End Road, London, El 4NS, UK; înstituto de Engenharia Biom6dica, Faculdade de Engenharia, Universidade do Porto, Praâ do Coronel Pacheco, N 1, 4050 PORTO, Portugal; Departments of Biomaterials^ and Periodontology*, Eastman Dental Institute, 256 Gray’s Inn Road, London, WCIX 8LD, UK

ABSTRAC T How cytometry (FCM) was used for evaluating the biological response of human cells to direct contact with materials with potential medical applications. The osteoblast-like cell line MG63 was grown on hydroxyapatite (HA) and P205-based glass reinforced hydroxyâtite composites (GRHA). Results suggested that the composites caused a delay in the progression of cells from GO/Gl into the S phase of the cell cycle. Moreover, although the cellular expression of collagen type I (COL I) was found to be similar in both materials, the level of osteocalcin (OC) was downregulated in the GR-HA compared with HA alone. The results of this study suggest that the FCM is a useful technique for assessing cell-biomaterial interactions. KEYWORDS : flow cytometry; cell culture; biocompatibility; hydroxyapatite INTRODUCTIO N A number of GR-HA materials with improved mechanical properties have been developed. However, their possible use for the repair and reconstruction of natural bone requires a critical biological evaluation. In vitrostudies of cytocompatibility have been carried out using several different assay methods, but the technique of flow cytometry (FCM) has previously not been widely used in thisfield.The present study has therefwe examined the use of the FCM technique to investigate cell-biomaterial interactions. More specifically, the effect of HA and GR-HA on the cell cycle and on the exiM-ession of COL I and OC were examined in a human osteosarcoma cell line. MATERIAL S AND METHOD S Preparation of discs A glass with a composition of 50.0%P2O5, 16.5%CaO and 33.5%MgO (mol %) was prepared. The composites were obtained by mixing 2.5 and 4.0 wt% of this glass with commercial hydroxyapatite. A detailed procedure for the preparation of this glass and composites has been reported previously [1]. The HA and the composites powders were uniaxially p-essed at 288MPa to form disc shaped samples. The discs were sintered at 1300 C for Ih, followed by natural cooling inside the furnace. All disc specimens were mechanically polished to 1 ^m finish, ultrasonically degreased and cleaned in ethanol followed by deionized water. Prior to cell culture the discs were sterilized in a dry atmosphere in a furnace at 180 ^C far 60 min. 575

576


Cell culture The MG63 cells were routinely cultured at 37 C in a humidified atmosphere of 5% CO2 in air, in 75cm^ flasks containing 10ml of alpha Minimum Essential Medium (a-MEM), 10% foetal calf serum (PCS), 2mM L-glutamine, 50IU/ml penicillin and 50 |ig/ml streptomycin. Media were change every third day and, for sub-culture, the cell monolayer was rinsed twice with phosphatebuffered saline (PBS) and incubated with trypsin-EDTA solution for 10 min at 37 ^C. The trypsin was inactivated by adding the complete medium at room temperature, the cells washed twice by centrifugation and resuspended in complete medium for re-seeding and growing in new culture flasks. The 30nmi discs were placed into 6-well culture plates. Aliquots of 500|il of cells were carefully placed onto the discs and allowed to settle for 2h in the incubator at 37 ^C, after which 2.5ml of complete medium was added. In control cultures the cells were placed directly into the plastic culture dishes in 2.5ml of medium, at the same density of cells per unit area as placed onto the discs. Cell cycle progression The cells were plated at 5000 cells/cm^ on the 30nun discs and the plastic dishes as described above. Approximately 36h after plating they were stained for DNA content using propidium iodide (PI), as follows. The cells were detached and suspended in 2ml of 70% ethanol for 30 min, then centrifuged and resuspended in 400|Lil of PBS. They were incubated with 100|Lil of RNase A (Img/ml) and 100|il of PI (400|il/ml) at 37 ^C for 45min and analyzed by PCM as described bellow. Immunofluorescent staining of antigens The cells were plated at 10000 cells/cm^ on 30nmi discs and directly on plastic dishes. The expression of collagen type I (COL I) and osteocalcin (OC) were measured after 6 days of culture. The cells were washed twice with PBS and detached using 20mM EDTA in PBS for approximately 5 min at 37 C. They were centrifuged at 400g for 7min and the pellet was resuspended and again centrifuged. They were fixed in l%(w/v) paraformaldehyde in PBS for 30 min, then washed by centrifugation and resuspended in washing buffer. This contained PBS, 2% PCS and 0.05% sodium azide. Aliquots of 10^ fixed cells were used to measure the level of each antigen by PCM. Cells were first permeabilized for 10 min using washing buffer with 0.1%(w/v) saponin, then washed, centrifuged and resuspended in washing buffer. Cells were incubated for 60 min at room temperature in either mouse monoclonal antibody (mAb) against human COL I, diluted 1:10, or in rabbit polyclonal antibody against human OC, diluted 1:100. Mouse IgGl and normal pre-inmiune rabbit serum were used as negative controls. Cells were then washed, centrifuged and resuspended again in washing buffer with 0.1% saponin. Secondary antibodies, fluorescein isothiocynate (PITC)conjugated rabbit anti-mouse IgG and FITC-conjugated swine anti-rabbit IgG, diluted 1:20, were added for 30 min at room temperature. Cells were washed again and resuspended in 400|xl of washing buffer and analyzed. FCM analysis The light scattering properties andfluorescenceof cells stained with PI and PTTC were measured on a PACScan flow cytometer (Becton Dickinson). Analysis were performed on 10000 cells. Data were collected, stored and analyzed with CELLQuest Software (Becton Dickinson).

Biological Evaluation of Glass ReinforcedHydroxy apatiteby Flow Cytometry:M.A. Lopes et al.

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RESULT S AND DISCUSSIO N The light scattering properties of FCM can be used to examine certain fundamental moq)hological characteristics of cells, since the intensity of light scattered at small angles (forward scatter; FSC) is considered proportional to cell size and the orthogonal light scatter (side scatter; SSC) is proportional to the granularity [2,3]. In the present experiments there were no significant differences in either FSC or SSC between cells grown on the GR-HA, HA or plastic dishes (data not shown). The cell cycle is the entire process by which a cell divides into two identical daughter cells. It is divided into five main phases: GO and Gl (pre-DNA synthesis; resting or interphase), S (DNA synthesis), G2 (post-DNA synthesis) and M (mitosis). The red fluorescence signals of cells stained with PI were used to obtain histograms of the DNA content of the individual cells in the MG63 cultures grown in the presence of the different materials, as described in the Materials and Methods. A representative histogram can be seen in Figure 1 in which the first peak corresponds to cells in GO/Gl, having a 2n DNA content, and the second peak corresponding to cells with G2+M DNA content of 4n. Cell scored in the trough have a DNA content intermediate between Gl and G2+M, and they usually considered to represent cells in S phase. As can be seen from the results shown in Figure 2, a higher proportion of quiescent cells with a lower proportion of cells in S phase was detected in the cultures grown on GR-HA compared with HA and control cells. This findings suggest that the GR-HA causes a delay in the progression of the MG63 cells through the cell cycle, although further studies are required to determine whether these materials affect the subsequent proliferation capacity of the cells. We also used the FCM to evaluate the effect of the GR-HA on cellular function, as measured in the present study by the levels of expression of COL I and OC, both of which have previously been shown to be key antigens in bone development, repair and regeneration. The results in Figure 3 show thefluorescencevalues of the GR-RA and HA cultures normalized to that of the cells grown directly on plastic culture tissue dishes. Although there was little difference in COL I expression between the different cultures, it is notable that OC levels were significantly lower in the GR-HA composites compared with HA. This may be explained at least partly by a delay in differentiation in the GR-HA cultures resulting from the apparently inhibitory effects of this material on the progression of the MG63 cells into the S phase of the cell cycle. Further studies were carried out to clarify the biological effects of the GR-HA materials. X-Ray diffraction (XRD) revealed a beta-tricalcium phosphate phase ((J-TCP) in addition to HA in all the composites. Table 1 shows the Rietveld analysis of these phases. Since it is known that P-TCP has a higher dissolution rate than HA, different cellular response would be expected. Moreover, we also found a slightly higher water contact angles of the GR-HA composites compared with HA (unpublished results), suggesting their higher hidrophobicity may also have an important influence on the progression of the cell cycle progression and hence on the functional activity of the cells.

Table 1 - Rietveld results for HA and GR-HA. HA

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Figure 2 - Frequency distribution of cell cycle phases at 36 h.

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Figure 3 - Expression levels of osteocalcin (OC) and collagen type I (COL I)

CONCLUSION The results of this study show that while the GR-HA materials appeared to have no major deleterious effects on MG63 cells, they nevertheless caused some perturbation in the progression of the cells through the cell cycle and also appeared to alter the regulation of a pivotal bone antigen. Moreover, our results show that the FCM technique is likely to be powerful new tool for assessing biocompatibility. ACKNOWLEDGMENTS The support of the JNICT (BD 1355 and project PBICT-C-CTM-1890-95) is gratefully acknowledged. REFERENCES [1] J. D. Santos, J. C. Knowles, F. J. Monteko, G. W. Hastings; Biomaterials, 15 (1994) 5. [2] M. G. Ormerod; How Cytometry, 1994, IRL Press. [3] H. M. Shapiro; Pratical Row Cytometry, 1995, Wiley-Liss Inc.


EVALUATIO N OF MACROPHAG E RESPONS E TO CERAMI C PARTICLE S BY FLOW CYTOMETRY : ANALYSI S OF PHAGOCYTOSI S AND CYTOTOXICIT Y I. Catelas\ R. Ma^chand^ L’H. Yahia\ O. L. Huk^ Êcole Polytechnique de Montreal, Institute of Biomedical Engineering, P.O. Box 6079, Succ. Centre-Ville, Montreal, Quebec, Canada, H3C-3A7, ^Microbiology and Immunology Department, Ste-Justine Hospital, Montreal, Canada, ^Lady Davis Institute for Medical Research, Jewish General Hospital, Montreal, Canada, ABSTRAC T Using the J774 cell line, we designed an in vitro model to analyse the effects of size, concentration and composition of ceramic particles (AI2O3 and Zr02) on phagocytosis and cell mortality by flow cytometry. The kinetic of phagocytosis was also analysed to determine at what time cell stimulation is maximal. Results reveal that phagocytosis increases with size and concentration for particles up to 2 |Lim. For larger particle range (up to 4.5 |im), phagocytosis reaches a plateau, which suggests a saturation of phagocytosis, most likely dependant on overall particle volume ingested. Cytotoxicity studies revealed that macrophage mortality increases with particle size and concentration for size greater than 2 |im and for concentrations up to 500 particles per macrophage. Smaller particles (0.6 |Lim) cause significant cell mortality only at higher concentrations (up to 1250 particles per cell), and the mortality is still very low (< 10%). No significant difference appears between AI2O3 and Zr02. Kinetic studies revealed that phagocytosis of the particles begins very early after cell exposure, increasing with time and particle concentration, and reaching a plateau at 15 hours, which implies that the optimum period to evaluate cellular response to particulate debris should be between 15 and 24 hours of incubation. KEYWORD S Ceramics, particles, macrophages, flow cytometry INTRODUCTIO N Aseptic loosening, initiated by wear particles, is without a doubt, the main cause of Total Hip Arthroplasty (THA) failure. Particle composition has been thought to play a role in the inflammatory reaction to wear debris generated from THAs. In this regard, ceramics have been considered bioinert. However, recent studies have shown that even ceramic particles can induce a macrophage foreign body response [1], and stimulate the production of bone resorbing mediators in vitro[2]. While analysis of pseudomembranes retrieved at revision surgery has provided much information on the particle sizes generated in vivo and has identified a myriad of inflammatory mediators that are released by macrophages, very little is understood about the exact parameters to which the macrophage responds and about the kinetics of this response. Indeed, many confounding factors are inherent in clinical studies, and it is difficult to determine the relative importance of each parameter, such as size, concentration and composition of the particles on ultimate macrophage response. Using flow cytometry, we developed a new in vitro model for 579

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testing cellular response [3] and the standardisation of flow cytometry for this type of in vitro model should be applicable to the study of cellular response to any kind of biomaterial. The purpose of this study was to analyse the effects of size, concentration and composition of ceramic particles (alumina AI2O3 and zirconia Zr02) on phagocytosis and cell mortality, using flow cytometry. The kinetics of phagocytosis was also analysed to determine at what time cell stimulation is maximal. MATERIAL S AND METHOD S Particles AI2O3 (Alcoa (Bauxite,Arkansas)), Alcan (Jonquieres,Quebec),Durmax (Chatellerault, France)) and Zr02 particles partially stabilised with yttrium (Unitec Ceramics (Stafford, England)) were commercially obtained in the phagocytosable range: 0.6 (Alcan), 1.3 (Alcoa), 2.4 (Alcan) and 4.5 jim (Durmax) for AI2O3 and 0.6 jim for Zr02 (Unitec Ceramics). Particles were sterilised by ethylene oxide. Cell culture The J774 cell line (ATCC, USA) was used in this in vitromodel. Macrophages were cultured and maintained in RPMI 1640 tissue culture medium (Gibco) with 10% foetal bovine serum and 2% antibiotics (gentamicin, penicillin, streptomicin). Cells were exposed to ceramic particles in tubes, containing 4.10^ cells in 1 ml of culture media with a particle suspension at concentrations varying from 30 to 1250 particles per macrophage. Culture tubes containing macrophages but no particles served as negative controls. Incubations were conducted at 37^C, in a 5% CO2 environment, for 24 hours. The tubes were then washed with PBS at pH 7.2, and cells were stained with propidium iodide (PI) for 4 minutes, at room temperature. Tubes were again washed and data were collected using flow cytometry. Methods Phagocytosis and cytotoxicity tests were conducted using flow cytometry (FACScan, Becton Dickinson). Flow cytometry analysis of phagocytosis determines the percentage of macrophages that have phagocytosed the particles (stimulated macrophages), by detecting changes in cell granularity and/or size following particle ingestion. This percentage is called the phagocytosi s index. The cytotoxic effect of particles is evaluated as cell mortality, and is obtained by measuring changes in cell fluorescence, after staining stimulated macrophages with propidium iodide. Kinetic studies allowed us to evaluate, when controlling for size, the progression of phagocytosis with time and concentration of particles. Kinetic tests (5 min to 24 hours) were conducted on AI2O3 at 1.3 jxm. RESULT S AND DISCUSSIO N Phagocytosis Figure 1 presents the effect of particle sizeand concentrationon phagocytosi s index. In the small particle range (< 2 |Lim), the phagocytosis index increases with particle size and concentration. In the larger particle range, the index reaches a maximum, independent of concentration. This suggests a saturation of phagocytosis, most likely dependant on particle volume ingested. As particle concentration increases, the percentage of stimulated macrophages increases. For particles larger than 2 |im, more than 80 % of macrophages are stimulated, independent of concentration. This later effect of maximal phagocytosis for large particles that was independent of concentration suggests that the larger the particle, the higher its chance of being seen by the cell and phagocytosed.

Evaluation of Macrophage Response to Ceramic Particles by Flow Cytometry: I. Catelas et al. 581 100 ^90

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Figure 2: Effect of particle concentration and composition on phagocytosis index

Figure 2 presents the effect of particle concentrationand composition(AI2O3 vs TxOi) when controlling for size at 0.6 |im. At that size, there is a direct relation between phagocytosis index and concentration, and no significant difference exists between AI2O3 and 2^-02. Figure 3 presents the studies of AI2O3 phagocytosis, when controlling for size at 1.3 |xm. Phagocytosis of the particles begins very early after exposure to the cells, which is similar to the time required to ingest bacteria for which macrophages have specific receptors [4]. Within 2 hours of incubation, there is a direct relation between phagocytosis index and particle concentration. After 15 hours of incubation, macrophage response to particle stimulation plateaus, which implies that the optimum period to evaluate cellular response to particulate debris should be between 15 and 24 hours of incubation. There is an initial artefactual decrease in phagocytosis from 0 to 2 hours which likely represents the adherence phenomenon of particles onto the macrophage membrane before particle internalisation begins.

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Figure 4: Effects of particle size and concentration on cell mortality

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Cytotoxicity Figure 4 presents the effects of AI2O3 particle size and concentration on cell mortality. Results reveal that macrophage mortality increases with particle size and concentration,and the later is only significant for particle size greater than 2 |im for concentrations up to 500 particles per macrophage. Smaller particles (0.6 |im) cause significant cell mortality only at higher concentrations (up to 1250 particles per cell), and the mortality is still very low (< 10%). However, macrophage stimulation was elevated at that size, as evidenced by the phagocytosis index as high as 80%. Despite evident stimulation at this combination of parameters, the cell is not sufficiently triggered to die. This demonstrates a distinction between the toxic and stimulatory effects of particulate debris on macrophages. With regards to composition, there was no difference in mortality between AI2O3 and Zr02 at the size (0.6 \xm)and concentrations we analysed. CONCLUSION S The evaluation of the macrophage inflammatory response in relation to various particulate debris parameters has not been thoroughly studied at the cellular level. Whereas microscopic evaluation of phagocytosis provides qualitative information, flow cytometry allows us to quantify the morphologic changes observed. The results of this study demonstrate that flow cytometry is a quantitative and objective technique to evaluate the inflammatory response to particulate debris. More specifically, we have shown that macrophage response to ceramic debris depends on size and concentration, which is in agreement with results in the literature [5] and is independent of ceramic composition. Our kinetic studies reveal that phagocytosis reaches a plateau at 15 hours of incubation, which implies that the optimum period to evaluate cellular response to particulate debris should be between 15 and 24 hours of incubation. Thus far, the clinical implications of this study indicate that if particulate debris in the size range up to 4.5 |Lim is generated in sufficient concentration, a macrophage response will be triggered, and that this response is most likely dependant on the volume of phagocytosed particles. Our results strongly support the concept that macrophages do not perceive absolute particle numbers, but rather, respond to overall particle volume. Future prosthetic design efforts should also focus on wear resistant interfaces to reduce debris. While not truly bioinert at the particulate level, the superiority of the ceramic-ceramic articulating interface in terms of wear resistance, merits continued study of this joint couple for THAs. REFERENCE S 1. Lerouge, S., Huk, O.L., Yahia, L’H., and Sedel, L., J. Biomed.Mater. Res. 1996, 32, 627633. 2. Nakashima, Y., Shuto, T Hayashi, K., Hotokebuchi, T., Yasuda, K., and Sugioka, Y., ORS, 1995, 780. 3. Catelas, I., Huk, O.L., Marchand, R., and Yahia L’H., In: BioceramicsVolume 9, Otsu, Japon 1996, 93-96. 4. Devrets, D.A., and Campell, P.A., J. of Immunol Methods1991,142, 31-38. 5. Shanbhag, A.R., Jacobs, J.J., Black, J., Galante, J.O., and Giant T.T., J. Biomed.Mater.Res. 1994,28,81-90.


STUD Y OF POROU S INTERCONNECTION S OF BIOCERAMI C ON CELLULA R REHABITATIO N IN VITRO AND IN VIVO J.X. Lu\ B. Flautre\ K. Anselme\ A. Gallur^, M. Descamps^ B. Thierry^ and P. Hardouin^ 1. IRMS, Institut Calot, Rue du Docteur Calot. 62608 Berck-Sur-MerFRANCE. 2. CRITT Ceramiques Fines, Z.I. du Champ de 1’Abbesse. 59600 MaubeugeFRANCE.

ABSTRAC T Our study would find the role of porous interconnections (PIC) on bone ingrowth and material degradation in tricalcium phosphate beta (P-TCP) with about 50% porosity, a size of 100-300 ^m macropores and of 30-100 îm PIC. In vitro,human osteoblasts were cultivated on discs with two delays: 14 and 28 days. In vivo,implants were implanted in middle diaphysis of both femurs with two delays: 12 and 24 weeks. The in vitro and in vivo samples were observed with histomorphometry (HMM) and scanning electronic microscopy (SEM). In vitroresults show that human osteoblasts penetrate sizes of PIC over 20 ^m diameter (dia.), set up and grow inside the bioceramic macropores. In vivo HMM results show that PIC directly influence bone ingrowth inside pores and material degradation. We notice that a PIC size over 50 \imdia. allows formation of new bone ingrowth inside the pores. The PIC density expresses the link between pores inside porous materials, assures the cells proliferation and tissular diflFerentiation by extra cellular and vascular exchanges. In resorbable materials, pores and PIC initial densities play a more important role than their sizes. KEYWORD : Bioceramic, porosity, porous interconnections, osteoblasts. INTRODUCTIO N The physico-chemical composition of tricalcium phosphate ceramics is identical to the mineral constituents of our skeleton. As far back as 1920, they were used as bone substitutes and showed the ability to be recolonized by bone tissue.^^^ Then, their use has been given up. But since 1970, a lot of studies investigated their behaviour in biological fluids. They were considered biocompatible either in vitroor in vivoP’^^ Meanwhile, the biological behaviour of these ceramics depends on factors involved in the processs such as the raw powders used, mineral phases, micro (pores size < 10 ^m) and macro (pores size > 100 ^m) porosity. Except for the chemical composition, degradation rate is directly influenced by the specific surface and the volume of microporosity; Macropores size and macroporosity volume influence bone ingrowth in these porous ceramics.^^’^^ Previous studies have mainly investigated bone ingrowth inside pores, but only a few ones have investigated the relationship between porous interconnections (PIC) and bone ingrowth. PIC is literally a pathway between pores to favour the cellular and vascular penetration which assures the bone ingrowth inside pores. The aim of our study was to observe the PIC influence on cellular and tissular bone ingrowth in porous materials. MATERIA L AND METHOD S Biomaterial: Beta tricalciimi phosphate (p-TCP) with a macroporosity of about 50 % and a microporosity of about 4.2% (measured by Hg porosimeter), a pores size of 100-300 îm and a PIC 583

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size of 30-100 ^m (measured by morphometry), and a Ca/P ratio of 1.55 – 0.03 (measured by Inductive Coupling Plasma) was used in the study. In vitrostudy: Osteoblasts were obtained from cancellous bone tissue of iliac crest of a 6-year-old boy. Cells were used after a second passage. p-TCP discs of 13.2 mm dia. and 1.4 – 0.2 mm height were sterilised by dry heat at 180 C during one hour. Discs were soaked in culture medium under vacuum. An osteoblasts suspension was prepared with culture medium in order to obtain 4 x 10^ cells/ml. One ml of cells suspension was then inoculated on the discs in each well. Complete medium was changed twice a week. After 14 and 28 days of incubation (37.0 – 0.2 C temperature, 5.0 – 0.1% CO2 and > 98% humidity), the discs were recuperated for HMM and SEM analysis. In vivostudy: p-TCP cylinders were implanted in rabbit. Implants size was 3 mm dia. and 6.1 – 0.2 mm length, and they were sterilised l^ 25 KGY gamma radiation. 8 white female New Zealand adult rabbits were 4 – 0.3 Kg body weight and 6-month-old before implantation. In rigorous asepsis conditions and under general anaesthesia, at the lateral external side of each femur a cavity of 3 mm dia. and 6 mm depth was made with an electric drill ((j) 2.8 mm, 200/minute speed) and then (() 3.0 mm manually in each middle femoral diaphysis. The created defect was washed from bone debris with saline solution. One implant was inserted in each cavity perpendicular to diaphysis. 17 and 2 days before sacrifice a double label of oxycycline and alizarin complexon was performed. After 12 and 24 weeks of implantation, the animals were sacrificed by an overdose of thiopental sodium (Nesdonalfi). The middle femoral shafts were removed and fixed in 10% neutral buffered formol for HMM analysis. Histomorphometry : Undecalcified bone preparation and polymethylmetacrylate embedding were used for each sample. Microradiographs were performed from 100 ^m thick sagittal sections. Then 50 ^m sections were stained with May Griinwald Giemsa staining for in vitrostudy and with Van Gieson’s Picro-Fuchsine staining for in vivo study. The following HMM parameters were measured: mineral apposition rate (MAR, mcm/d), new bone volume (BV/IV, %), residual material volume (MV/IV, %), material degradation rate (MDR, % = 100% - measured MV/IV / initial MV/IV), relative osteoid tissue volimie (OV/BV, %), macropores (MP) size (^m) and density (number/mm^), PIC size (^mi) and density (nimiber/mm^). Statistical analysis: Results were expressed by means and standard deviations. Impaired bilateral Mest (SYSTATfi) was used to compare the delays. Correlations (PIC mean size, MP mean size, PIC mean density, MP mean density and the BV/IV) were studied with Pearson correlation coefficient (r). RESULT S In vitrostudy: SEM analysis (A7 = 2): a cellular layer spread on the materials area, and covers the opened MP. On a broken surface of cylinders, osteoblasts are observed in the MP. After 28 days of culture the cells penetration is deeper than after 14 days. Some osteoblasts inside pores go through the PIC (Fig. 1). Nevertheless, no modification of the microstructure of the ceramic was observed after culture with SEM. On stained sections observed with transmitted light microscope (n = 4), cells are seen inside of PIC. The size of PIC containing is between 19.15 and 215.43 ^m dia. with a mean size of 78.67 – 37.41 îm. The maximalfrequencyPIC size with cells is 60 ^m dia. after 28 days of culture. In vivo study: MAR is higher in receiving bone near the implant than inside the implant. Inside pTCP, MAR and BV/IV increases (p < 0.05), and OV/BV and MV/IV decreases {p < 0.05) with time (Table 1). On stained sections, the pathways size with bone tissue is 19.15 - 201.06 ^m dia..

Study of PIC of Bioceramics on Cellular RehabitationIn Vitro and In Vivo: J.X. Lu et al.

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but no significant difference (the mean value of size mean of each sample and their standard deviation: 74.40 – 5.29 for T12 weeks and 77.32 – 5.47 for T24 weeks) is observed between the delays. Table 1.The HMM parameters before and after Implantation Delay n Bone P-TCP Implant Weeks MAR(mcm/d)MAR(mcm/d) BV/IV(%) OV/BV(%) MV/IV(%) MDR (%) TO 8 59.79 –3.23 12 8 1.78 + 0.32 0.91+0.17 35.00 + 6.64 7.48 + 5.16 52.89 + 4.58* 11.49 –7.66 24 8 1.7410.2 1 1.25 –0.10* 52.07+7.53* 3.42 + 2.93* 42.57 –7.71* 28.77 –12.90* * Compare between the delays with/? < 0.05. In microradiography, MP size and density increase with time. The curve of PIC distribution shows a translation in the biggest sizes with aflattenedcurve (Fig. 2). Mean size and density of PIC have increased (p < 0.05) with time, but no significant difference is observed for the PIC density between 12 and 24 weeks (Table 2). Figure 1. SEM in vitro p-TCP, 28 days: osteoblasts (short arrows) go througji the PIC (lc»ig arrow).

Figure 2. Porous Interconnections Distribution of p-TCP

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Table2. Characterisatio n of (3-TCP microstructur e before and after implantation n Delay Sample Macropores Porous Interconnection s n Density {nlvam^) (Weeks) Size (îm) Density (w/mm^) Size (fim) TO 8 14.46 –1.67 165.62 –6.60 26.45 –1.90 63.93 –3.81 12 8 17.58 –1.11* 178.15 –11.56 * 21.72–2.15 * 78.59–5.21 * 24 8 18.74 –2.49 202.63 – 12.20 * 18.41 –1.77* 93.83 –7.63*

D The mean value of size inean(or density meam) of every sample and their standard deviation. * Comparison between delays with/? < 0.05. Correlations: A correlation between initial MP size and density is observed (r = -0.789, p < 0.05). No correlation is foxind between the initial PIC size and density, neither between the PIC and the MP. At T12 weeks, a conelation between PIC density and MP density is seen (r = 0.765, p < 0.05), also a correlation between BV/IV and the PIC size (r = 0.733, p < 0.05), or between BV/IV and the pores density (r = -0.736, p < 0.05). At T24 weeks, there is also a correlation between the PIC size and MP size (r = 0.880, p < 0.01) or between the PIC size and MP density (r = -0.786, p < 0.05). DISCUSSION GALOIS et al.^^^ did not found rats dermfibroblastsinside HA (hydroxyapatite) or p-TCP implants with different pores size after 5 and 40 days of culture. However, our in vitrostudy shown that human osteoblasts have spread in the pores of porous ceramics and are able to through inside the

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PIC. However the osteoblasts are seen only in the first 3 rings of ceramic (1.2 – 0.2 mm depth) but never in the center. To avoid the hydrophobic phenomenon of porous ceramics, the air bubbles were extracted with vacuum and stained coloration was justified. According to our results, the osteoblasts can cross through 20 fim PIC size. KLAWITTER and HULBERT^^^ have noted that the PIC size of calcium aluminate should be at least 100 ^m size to favour the mineralised bone tissue growth, between 40 and 100 |iim for osteoid tissue and at last 5 - 15 ^m for fibrous tissue. SHIMAYAKI and MOONEY^^^ have studied comparatively 230 and 600 ^m dia. of pores with respectively 190 and 260 ^m dia. of PIC in porous HA implanted in rabbit femurs. Results shown a higher bone ingrowth for 600 ^m than 230 ^m pores size. In our study, MAR and BV/IV increase significantly with time. PIC size and density increase in time and facilitate the pathway of vessels. Mean size of PIC with mineralised bone tissue inside was about 70 ^m diameter. We have measured the Volkman canals size inside the ceramic on microradiography. Their size was 10-20 \\mdia. with a maximal frequency of 20 ^m size. It could be the reason why a correlation is observed between new bone formation and PIC size or and pores density after 12 weeks. Because of an increase of p-TCP degradation at 24 weeks no correlation with new bone formation is observed at this time, but close correlation yet in their microstructure. WHITE and SHORS^^^ considered that PIC size must be superior than 100 ^m. But, our results indicate that a PIC size greater than 50 ^m allows the mineralised bone formation inside pores of material. Because our two-dimensions measurement we could precize the maximal PIC size, but we failed to precize the minimal PIC size. CONCLUSIO N These results show that in vitro,human osteoblasts are able to cross through PIC and to spread inside the pores. A minimal PIC size of 20 fim dia. is required but the ideal size is about 60 ^m diameter. In vivo,a PIC size greater than 50 fim allows formation of new mineralised bone tissue. Density and size of PIC and MP can be modified by the biodegradation of the (3-TCP. New PIC can improve the cells, vessel and biological liquids penetration and facilitate the material degradation and bone ingrowth inside the material. Therefore for resorbable materials the MP and PIC initial density play a more important role than their size. REFERENCE S 1. F.H.ALBEE,^«w.5’wrg. 71(1920)32. 2. L. Galois, D. Mainard, K. Bordji, H. Membre, L. Marchal, B. Foliquet, D. Clement and J.P. Delagoutte, in "Actualitesen materiau"(Romillat, Paris,), 3 (1990) 361. 3. J.J. Klawitter and S.F. Hulbert, J. BiomedMater.Res.Symp.2 (1971) 161. 4. K. Shimazaki and V. Mooney, J. Orthop.Res.3 (1985) 301. 5. E. White andE.C. Shors, Dent Clin. Nor. America.30 (1986) 49. 6. D. Groot, in the Materialcharacteristicsversusin vivo behavior,P. Ducheyne and J. Lemons (eds.), N.Y. Acad. Sci., 523 (1988) 227. 7. R.Z.L. Geros, JR. Parsons, G. Daculsi, F. Drissens, D. Lee, ST. Liu, D. Peterson and M. in vivo behavior,P. Ducheyne and J. Lemons Walker, in the Materialcharacteristics versus (eds.), N.Y. Acad. Sci., 523 (1988) 268.


REPAI R OF OSTEOCHONDRA L DEFEC T USIN G ARTIFICIA L CARTILAG E

ARTICULA R

M. Hasegawa, A. Sudo, Y. Shikinami* and A. Uchida Department of Orthopedic Surgery, Mie University, Faculty of Medicine, 2 -174 Edobashi, Tsu City, Mie 514, and *Takiron CO., LTD., 405, Nagano, Yasutomicho, Shisou-gun, Hyogo 671-24, Japan

ABSTRAC T A new artificial articular cartilage was used for to repair an osteochondral defect. To test its biocompatibility, a large full-thickness defect was created in the patellar groove of a rabbit, and was filled with artificial articular cartilage tightly. Macroscopically, the repair tissues appeared as glistening smooth surface partially covering the implants. On histological and immunohistochemical evaluation, the ingrowth of bone and hyaline cartilage-hke tissue was found surrounding and within the implants. All the repairs with or without implants were associated with slight synovitis of the knee joint. The artificial articular cartilage [three-dimensional fabric (3-DF)] has bulk and surface biocompatibility, and could serve as both a scaffold for cartilage formation and a prosthesis. KE Y WORD S artificial articular cartilage, osteochondral defect, histological evaluation INTRODUCTIO N It has been reported that large articular cartilage defects of mature animals have little capacity for repair. Repair of full-thickness osteochondral defects has been attempted by various methods. Some experiments have achieved successful repair [1-7], but there is no highly acceptable method for complete repair of hyaline cartilage. The purpose of this study was to test the biocompatibility and repair potential of a threedimensional fabric (3-DF) developed as a new implant for artificial articular cartilage [8] to enhance the healing of large osteochondral defects in the rabbit knee. MATERIAL S AN D METHOD S The 3-DF used in this experiment is a prototype of FABRICUBE^^ (Takiron Co., Ltd). This triaxial-three dimensional (3A-3D) fabric was fabricated by the method described below. Low density linear polyethylene (LLDPE) was melted and used to the surface of ultra-high molecular weight polyethylene (UHMWPE) fibers that was a yam of 500 denier (50 filaments) to prepare a yarn with an average diameter of 400|im (Yam A). Yam A was treated with a hand operation practice machine which was designed to produce a 3 A-3D structure. The 3-DF is a block with an orthogonal fiber alignment and a size 10 x 10 x 3mm (Figure 1). Its compression behavior (mechanical strength and profile) was quite 587

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similar to that of natural cartilage. The surface of Yam A was oxidized by corona-discharge. Thereafter, the LLDPE was softened by heating, and a micro powder of unsintered hydroxy apatite was sprayed onto the surface. After cleaning the surface with water, partial exposure of the powder from the surface was confirmed. Fourteen skeletally mature female Jânese white rabbits, weighing 3.2 - 3.9 kg, were used After shaving, anteromedial arthrotomy was performed under intravenous pentobarbital sodium anesthesia (20mg/kg body weight). The patella was dislocated laterally, and the articular surface of the distal femur was exposed A full-thickness osteochondral defect, measuring 10mm from proximal to distal transversing the entire width of the patellar groove and 3mm in depth penetrating the subchondral bone plate generated in the patellar groove of the femoral condyle using osteotome, rongeur and curette. The defect on one knee was filled with sterile artificial cartilage with the surface of the implant was parallel to the original cartilage surface (group A; n= 14). The defect on the other knee was left empty as a control (group B; n = 14). Joint capsule and skin were sutured as separate layers. After the operation, all rabbits were allowed free cage activity without immobilization. Two rabbits with artificial articular cartilage on one side and control on the other were excluded because one died 11 days after the operation, and one sustained a dislocation of the patella on the knee with the implanted artificial cartilage. The remaining the 12 animals were sacrificed with a lethal dose of pentobarbital sodium administered intravenously from 2 to 12 weeks postoperatively. The distal part of the femur and synovium were examined macroscopically. For microscopical evaluation, the distal femur was removed, fixed in 10% buffered formalin, decalcified in Plank-Richlo solution at A C and embedded in paraffin. Sagittal sections 5-10|im thick were cut and stained with hematoxylin andeosin (HE) and alcian-blue. Immunostaining with an antibody to type n collagen was visualized by the avidin-biotin immunoperoxidase method Synovial tissue samples were collected from the suprâtellar pouch, fixed and embedded in the same way, then stained with HE to evaluate inflammation and immune responses.

Figure 1. artificial (FABRICUBE ).

articular

cartilage

Figure 2. Photomicrograph showing the ingrowth of cartilage into the implant at 8 weeks postoperatively (alcian-blue, original magnification x 100).

Repair of OsteochondralDefect Using Artificial Cartilage: M. Hasegawa et al.

589

RESULTS Macroscopic Observations Neither joint contracture nor infection was found in any animals. Slight intra-articular effusion was observed only at 2 weeks postoperatively in both groups, but decreased with time. In group A, the implant was covered with white glistening tissue on the medial and lateral side 2 weeks postoperatively. Tissues around the implant became thicker with time, but no thick tissues were formed on the center of the artificial cartilage by 12 weeks postoperatively. In group B, the defect was filled with soft brown tissue, and the surface was very rough and depressed below the surrounding original articular surface 2 weeks postoperatively. The defect was filled with white and opaque tissue, which did not reach the level of the original articular surface during the experimental period The synovial membrane was slightly hyperemic and hyperplasic in both groups 2 weeks postoperatively, but these observarions became more normal with time. Microscopic Observations In group A, cartilaginous rissue was found at the base and side of the implanted material at 2 weeks postoperarively. Ingrowth of fibrous tissue into the implant was seen, and the newly formed tissue on top of the implant was fibrous. The ingrowth of bone and cartilage into the implant was seen 4 weeks postoperatively (Figure 2). Hyahne cartilage-like tissue and bone were observed covering part of the implant in 8-week specimens (Figure 3). Hyaline cartilage-like tissue was stained by alcian-blue and immunostained by the antibody to type II collagen. Most cartilage in contact with the implant was replaced by bone. Abundant bone ingrowth into the implant was seen in the deep region 8 weeks postoperatively (Figure 4). No bone was found at the articular surface by 12 weeks postoperatively. No inflammatory reaction was observed In group B, the surface of the defect was depressed below the normal level by 12 weeks postoperatively. Fibrous tissue including scant hyaline cartilage-like tissue was seen in the superficial region and bone was observed in the deep region. Synovial rissue showed mild inflammation including lymphocyte infiltration and subsynovial fibrosis at 2 weeks in group A. However, these changes tended to diminish with time, and the same findings were seen in group B. No foreign body giant cells or debris particles were found.

Figure 3. Photomicrograph showing neocartilage formation covering part of the implant at 8 weeks postoperarively (HE, original magnification x 40).

Figure 4. Photomicrograph showing the ingrowth of bone into the implant at 8 weeks postoperatively (HE, original magnificarion x 40).

590


DISCUSSIO N Various methods for to repair of osteochondral defects have been reported using organic [1-3] and inorganic materials such as collagen matrix [4], carbon fiber [5], polytetrafluoroethylene [6], polyester [6] and polylactic acid matrix [7]. Artificial materials can have the mechanical function at an early stage and be easily shaped to fit the defect, but have drawbacks such as fixation and the risk of debris particle generation. When carbon fiber was used as a scaffold for repair of full-thickness osteochondral defects, cartilage repair was not obtained and minimal synovitis with pigmentation were observed [5]. Implantation of polytetrafluoroethylene (Tefron) and polyester (Dacron) showed early neocartilage formation without normal cartilage morphology and synovitis [6]. When polylactic acid matrix was implanted in cartilage defect with and without periosteal grafting, polylactic acid matrix with periosteal grafting showed formation of cartilage resembling articular cartilage [7]. The experiments described above involved small defects (3.0-3.7mm in diameter) produced in the knees of rabbits. Small osteochondral defects less than about 3mm in diameter have been suggested to be repaired naturally, although repaired cartilage never fully resembles normal hyaline cartilage biochemically [9]. In this study, we found early formation of cartilage in contact with the artificial cartilage in the osteochondral defects. The implant coated by hydroxy apatite was fixed firmly to the subchondral bone which is necessary for artificial cartilage. Inflammatory reactions against the implant were slight, shown as synovial hyperplasia in 2-week histological specimens. Complete repair of articular cartilage defects was not achieved during the experimental period because of large the size of the defect. However, this implant could be used as a prosthesis. We conclude that this implant has good bulk and surface biocompatibility, and could serve as both a scaffold for bone and cartilage formation and a prosthesis.

REFERENCES 1. 2. 3. 4. 5. 6. 7. 8. 9.

O’Driscoll, S. W., Keeley,F. W. and Salter, R. B. J. Bone JointSurg. 1986, 68-A, 1017-1034. Billings, Jr., E., von Schroeder,H. P., Mai, M. T., Aratow, M., Amiel, D., Woo, S. L.-Y. and Coutts, R. D. Acta OrthopScand.1990, 6 1, 201-206. Grande, D. A., Pitman, M. I., Peterson, L., Menche,D. and Klein, M. J. Orthop.Res. 1989, 7, 208-218. Speer, D. P., Chvapil, M., Volz, R. G. and Holmes, M. D. Clin. Orthop.1978, 25, 326-335. Muckle, D. S. and Minns, R. J. J. Bone Joint Surg. 1990, 72-B, 60-62. Messner, K. and Gillquist, J. Biomaterials1993, 14, 513-521 Von Schroeder, H. P., Kwan, M., Amiel, D. and Coutts, R. D. J. Biomed.Mater.Res. 1991, 2 5, 329-339. Shikinami, Y. and Kawarada, H. Biomaterials(submitted). Furukawa, T., Eyre, D. R., Koide, S. and Glimcher, M. J. J Bone Joint Surg. 1980, 62.A, 79-89.


CALCIUM PHOSPHATE CERAMICS AS CONTROLLED RELEASE SYSTEMS FOR FGF-2 Midy v.* +, Hollande E.+, Rey C*. and Dard M§. * Laboratoire des Mat^riaux et Interfaces. Phosphates et Biomat^riaux, UPRESA CNRS 5071, 38 rue des 36 Ponts, 31400 Toulouse, France ^Laboratoire de Biologie Cellulaire, 38 rue des 36 Ponts, 31400 Toulouse, France § Merck Biomaterial Research, Frankfurter Str. 250,64271 Darmstadt, Germany KEY-WORDS : Calcium phosphate ceramics, Fibroblastic Growth Factor-2, Adsorption. ABSTRAC T The aim of our study was to determine the capacity of calcium phosphate ceramics to serve as carriers for Fibroblastic Growth Factor-2 (FGF-2), which is able to locally stimulate bone formation in vivo. Two different apatitic substrates were tested: stoichiometric hydroxyapatite (HA), widely used as an osteoconductive biomaterial, and nanocrystalline carbonated apatite (CA), very close to bone mineral crystals. Adsorption of the growth factor was quantified by ^^Î and reached 85% of the initial concentration for CA, and 10%, for HA. Release of FGF-2 was about 20%, of the adsorbed amounts for both compounds. In fetal bovine aortic endothelial cell culture, the FGF-2 released showed a strong loss of bioactivity possibly due to the presence of ceramic particles and/or to conformational changes. INTRODUCTIO N Bioactivity is an essential characteristic of several calcium phosphate materials. It involves four steps: surface modification of the ceramic, nucleation of carbonate apatite on the surface of the material from supersaturated body fluid, adsorption of protein, and adhesion and differentiation of bone cells. Althought calcium phosphate materials generally used as biomaterials differ substantially from the mineral part of bone, the precipitation, on their surface, of carbonated apatite crystals very analogous to bone mineral crystals is believed to play a dominant part in the biological response. The relationship between this neoformed layer and bone proteins or growth factors and their action on bone cells however remains largely unknown [1]. It is widely accepted that a multitude of factors, from systemic hormones to local regulatory factors, cytokines, and prostaglandins, act together to regulate the coupling between bone formation and bone resorption. Fibroblastic Growth Factor-2 (FGF-2) is one of the growth factors which can locally stimulate bone formation in vivo [2]. It is produced by osteoblasts and stored in skeletal tissues, and it has been shown to stimulate proliferation of osteoblasts, chondrocytes and periosteal cells [3,4,5]. In vivo studies with FGF-2 also suggest a potential therapeutic role in the treatment of bone loss [6]. Based on these findings, it is speculated that the association of FGF-2 with calcium phosphate ceramics may be useful to facilitate fracture repair. The aim of this report is to quantify the adsorption and release of FGF-2 from two different apatitic substrates, stoichiometric hydroxyapatite (HA) and nanocrystalline carbonated apatite (CA). In additioin the activity of the 591

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FGF-2 released was evaluated. This last point appears particularly important to assess the ability of the calcium phosphate ceramics to act as controlled release systems. MATERIA L AND METHOD S Hydroxyapatite (HA) powder was obtained from Merck KGaA (Darmstadt, Germany), carbonated apatite (CA) powder was prepared as previously described [7]. Recombinant human FGF-2 obtained from Merck Biomaterials (Darmstadt. Germany), was added to the ceramics at various concentrations (0.5-1-2.5-5 Mg/ml). The growth factor in the bound and in the unbound fractions was evaluated by ^^Î counting. The amount of FGF-2 released in a Tris Buffer Salt (TBS) solution was also determined by using ^^î labeled FGF-2. The in vitroactivity of FGF-2 released after adsorption, from the calcium phosphate ceramics into TBS buffer was determined by using proliferation assay on FBAE (foetal bovine aortic endothelial) cells and on osteoblast-like cells. At the end of the period, the cells were counted using a Coulter Counter and the amount of active FGF-2 was determined by reference to a previously obtained standard curve. RESULTS AND DISCUSSION CA is a calcium-deficient, poorly crystalline apatite. Infra-red spectroscopy showed labile, non-apatitic carbonate and phosphate enviromnents probably located at the surface of the crystals [8], These characteristics are very similar to those of mature human bone mineral. Figure 1 shows the adsorption isotherm of FGF-2 on HA and CA. The adsorption of FGF2 reached 85% of the amount contained in solution for CA, and about 10% for HA. The experimental conditions did not allow saturation of the adsorption sites.

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Bioceramics Volume 10 (Biocermaics)

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