Titanium Powder Metallurgy: Science, Technology and Applications [1 ed.] 0128000546, 9780128000540

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Titanium in Medical and Dental Applications

Related titles Titanium Powder Metallurgy: Science, Technology and Application (ISBN: 978-0-12-800054-0) Titanium Alloys: Modelling, Microstructure, Properties and Applications (ISBN: 978-1-84569-375-6) Metals for Biomedical Devices (ISBN: 978-1-84569-434-0) Precious Metals for Biomedical Applications (ISBN: 978-0-85709-434-6) Materials Science for Dentistry (ISBN: 978-1-84569-529-3)

Woodhead Publishing Series in Biomaterials

Titanium in Medical and Dental Applications

Edited by

Francis H. Froes Ma Qian

An imprint of Elsevier

Woodhead Publishing is an imprint of Elsevier The Officers’ Mess Business Centre, Royston Road, Duxford, CB22 4QH, United Kingdom 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States The Boulevard, Langford Lane, Kidlington, OX5 1GB, United Kingdom © 2018 Elsevier Inc. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www.elsevier.com/permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notices Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers must always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library ISBN: 978-0-12-812456-7 (print) ISBN: 978-0-12-812457-4 (online) For information on all Woodhead publications visit our website at https://www.elsevier.com/books-and-journals

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List of contributors

S. Akyol University of Istanbul, Istanbul, Turkey A.Y. Arbenin St. Petersburg State University, St. Petersburg, Russia J.D. Avila Washington State University, Pullman, WA, United States M. B€ aker Technische Universit€at Braunschweig, Institute for Materials, Braunschweig, Germany A. Bandyopadhyay Washington State University, Pullman, WA, United States K. Bartha Charles University, Prague, Czech Republic C. Battocchio Roma Tre University of Rome, Rome, Italy H.S. Baumgarten Penn Dental Medicine, Philadelphia, PA, United States A. Bendavid Commonwealth Scientific and Industrial Research Organisation, Lindfield, NSW, Australia B. Ben-Nissan University of Technology Sydney, Ultimo, NSW, Australia S. Bose Washington State University, Pullman, WA, United States F. Brunke Technische Universit€at Braunschweig, Institute for Materials, Braunschweig, Germany A.H. Choi University of Technology Sydney, Ultimo, NSW, Australia  ´zˇek Charles University, Prague, Czech Republic J. Cı C.M. Cotrut Universisty Politehnica of Bucharest, Bucharest, Romania M. Dinu National Institute for Optoelectronics, Magurele, Romania L. Dluhosˇ Timplant s.r.o., Sjednocenı´, Ostrava, –Polanka, Czech Republic T. Duerig Confluent Medical Technologies, Fremont, CA, United States

xvi

List of contributors

T. Ebel Helmholtz-Zentrum Geesthacht, Geesthacht, Germany Y. Estrin Monash University, Clayton, VIC, Australia; NUST MISiS, Moscow, Russia S. Franchi Roma Tre University of Rome, Rome, Italy J.L. Fraysse Forecreu SAS, Malicorne, France F.H. (Sam) Froes Consultant to the Titanium Industry, Tacoma, WA, United States S. Gatina Ufa State Aviation Technical University, Ufa, Russia P. Gubbi Zimmer Biomet Dental, Palm Beach Gardens, FL, United States T. Hanawa Tokyo Medical and Dental University, Tokyo, Japan Y.L. Hao Institute of Metal Research, Chinese Academy of Sciences, Shenyang, China Y. Hao Shenyang National Laboratory for Materials Science, Institute of Metal Research, Chinese Academy of Sciences, Shenyang, China P. Harcuba Charles University, Prague, Czech Republic W.T. Hou Institute of Metal Research, Chinese Academy of Sciences, Shenyang, China G. Iucci Roma Tre University of Rome, Rome, Italy E. Ivanova Swinburne University, Hawthorn, VIC, Australia M. Janecˇek Charles University, Prague, Czech Republic L. Jia Northwest Institute for Nonferrous Metal Research, Xi’an, China C. Kasper University of Natural Resources and Life Sciences, Vienna, Austria K. Kondoh Osaka University, Osaka, Japan A. Kumar University of Texas at El Paso, El Paso, TX, United States M. Landa Institute of Thermomechanics, Academy of Sciences, Prague, Czech Republic

List of contributors

xvii

R. Lapovok Deakin University, Geelong, VIC, Australia M. Leary Centre for Additive Manufacturing, School of Engineering, RMIT University, Melbourne, VIC, Australia S.J. Li Institute of Metal Research, Chinese Academy of Sciences, Shenyang, China S. Li Shenyang National Laboratory for Materials Science, Institute of Metal Research, Chinese Academy of Sciences, Shenyang, China N. Liu Northwest Institute for Nonferrous Metal Research, Xi’an, China T.C. Lowe Colorado School of Mines, Golden, CO, United States A.E. Medvedev Monash University, Clayton, VIC, Australia R.D.K. Misra University of Texas at El Paso, El Paso, TX, United States M. Morinaga Toyota Physical and Chemical Research Institute, Nagoya, Japan M. Niinomi Tohoku University, Sendai, Japan; Osaka University, Osaka, Japan; Nagoya University, Nagoya, Japan; Meijyo University, Nagoya, Japan F. Ory Forecreu SAS, Malicorne, France E.V. Parfenov Ufa State Aviation Technical University, Ufa, Russia V. Polyakova Ufa State Aviation Technical University, Ufa, Russia D. Preisler Charles University, Prague, Czech Republic V. Pruna Institute of Cellular Biology and Pathology “Nicolae Simionescu” of the Romanian Academy, Bucharest, Romania M. Qian RMIT University, Melbourne, VIC, Australia A. Ramezannejad RMIT University, Melbourne, VIC, Australia I. Sabirov IMDEA Materials Institute, Getafe, Madrid, Spain M. Santi Roma Tre University of Rome, Rome, Italy J.J. Scutti Boston Scientific, Raynham, MA, United States

xviii

List of contributors

V. Secchi Roma Tre University of Rome, Rome, Italy I. Semenova Ufa State Aviation Technical University, Ufa, Russia T.B. Sercombe University of Western Australia, Crawley, Australia H. Sibum Ingenieurb€ uro f€ ur Werkstoffverarbeitung und -design, Grevenbroich, Germany C. Siemers Technische Universit€at Braunschweig, Institute for Materials, Braunschweig, Germany V.M. Smirnov St. Petersburg State University, St. Petersburg, Russia R. Soba Terumo Corporation, Kanagawa, Japan D.B. Spenciner Rensselaer Polytechnic Institute, Raynham, MA, United States J. Stra´sky´ Charles University, Prague, Czech Republic Y. Tanabe Terumo Corporation, Kanagawa, Japan H.P. Tang Northwest Institute for Nonferrous Metal Research; Xi’an Sailong Metal Materials Co. Ltd., Xi’an, China R.G. Thompson Anatomics Pty. Ltd., St Kilda, VIC, Australia I. Titorencu Institute of Cellular Biology and Pathology “Nicolae Simionescu” of the Romanian Academy, Bucharest, Romania J. Umeda Osaka University, Osaka, Japan R.Z. Valiev St. Petersburg State University, St. Petersburg; Ufa State Aviation Technical University, Ufa, Russia A. Vladescu National Institute for Optoelectronics, Magurele, Romania T. Wojtisek Zimmer Biomet Dental, Palm Beach Gardens, FL, United States D. Wolter Osteosynthese-Institut GbR, Ahrensburg, Germany C.S. Xiang Xi’an Sailong Metal Materials Co. Ltd., Xi’an, China Q.S. Xu Institute of Metal Research, Chinese Academy of Sciences, Shenyang, China

List of contributors

xix

W. Xu RMIT University, Melbourne, VIC; Macquarie University, Sydney, NSW, Australia R. Yang Institute of Metal Research, Chinese Academy of Sciences, Shenyang, China E.G. Zemtsova St. Petersburg State University, St. Petersburg, Russia L.-C. Zhang Edith Cowan University, Perth, Australia P. Zhao Xi’an Sailong Metal Materials Co. Ltd., Xi’an, China S. Zhao Institute of Metal Research, Chinese Academy of Sciences, Shenyang, China

Preface

Since the 1950s, titanium has been used in orthopedic applications (orthopedics is the branch of surgery concerned with conditions involving the musculoskeletal system). Orthopedic surgeons use both surgical and nonsurgical means to treat musculoskeletal trauma, spine diseases, sports injuries, degenerative diseases, infections, tumors, and congenital disorders. Titanium alloys are the standard material of choice for orthopedic devices such as hip joints, bone screws, knee joints, spinal fusion cages, shoulder and elbow joints, bone plates, and scaffolds. Titanium has been selected as a metal of choice in the orthopedic segment of the market in part because of its general corrosion resistance. In fact, titanium is inert in the human body, and has proven to be strong, fracture resistant, and compatible with bone density, has a low modulus, and displays nonmagnetic characteristics (titanium alloys offer a very low quantity of artifacts in magnetic resonance imaging, or MRI), making it an excellent material of choice in the orthopedic arena. Titanium is also used in cardiovascular applications which relate to the circulatory system (involving the heart and blood vessels which carry nutrients and oxygen to the body tissues and remove carbon dioxide and other wastes). Cardiovascular diseases are conditions which affect the heart and blood vessels including arteriosclerosis, coronary artery problems, heart valve concerns, arrhythmia, heart stoppage, hypertension, orthostatic hypotension, and congenital heart disease. Titanium’s advantages for cardiovascular applications are similar to those listed above for orthopedic applications. This book presents papers which summarize the advantages provided by titanium in medical and dental applications. A number of body implants fabricated from titanium are discussed. Dental uses of titanium are presented, with many of the advantages for this application being similar to the characteristics presented for medical use. Titanium’s use in surgical instruments (its low density, translating to reduced weight, being an advantage here) wheel chairs, walkers, and walking sticks is featured. This is followed by papers discussing the specifications for titanium for implants and a summary of the mechanical properties of titanium alloys suitable for implants (with emphasis on the Ti-6Al-4 V alloy and commercially pure grades of titanium). Articles on other titanium-based alloys such as nitinol (shape memory) are then documented for both body and dental applications. World-wide, more than 1000 tons (2.2 million pounds) of devices made from titanium are implanted in patients every year. Requirements for joint replacement

xxii

Preface

increase as people live longer or damage themselves through hard sports play or jogging, or sustain critical injuries due to traffic and other accidents. We thank all the contributors to this book and hope that you, the reader, enjoy what we have put together. Francis H. (Sam) Froes Consultant to the Titanium Industry, Tacoma, WA, United States Ma Qian RMIT University, Melbourne, VIC, Australia

About the editors

Professor Francis H. (Sam) Froes has been involved in all aspects of the titanium field for more than 40 years. After receiving a BSc from Liverpool University and an MSc and PhD from Sheffield University, all in Physical Metallurgy, he was employed by a primary titanium producer, Crucible Steel Company of Pittsburgh, Pennsylvania. At this organization, he was leader of the titanium group involved in research, technology, applications, marketing, and sales of titanium. He then spent time at the USAF Materials Laboratory in Dayton, Ohio, where he was supervisor of the light metals group (which included titanium). During this time, he was the keynote speaker at the first International Titanium Association Conference in 1984 held in Keystone, Colorado. This was followed by 17 years at the University of Idaho, Moscow, Idaho, where he was an Institute Director (Materials and Advanced Processes, IMAP) and Department Head of the Materials Science and Engineering Department. Here, he again emphasized research and development on titanium and its alloys. During this time, he was the Chairman of the Seventh World Titanium Conference in San Diego, California, in 1992. Subsequently, Dr. Froes has retired to Tacoma, Washington, where he continues to write papers and edit books on all aspects of titanium. He has more than 850 publications and in excess of 60 patents, the majority on various aspects of titanium. Since the early 1980s, he has taught the ASM International short course on “Titanium and its Alloys.” He has organized more than 15 symposia on various aspects of titanium science, technology, and applications, including in recent years four co-sponsorships of sessions on cost-effective titanium. His recent interests have included powder metallurgy in general, with emphasis on additive manufacturing, and medical and dental applications of titanium. Dr. Froes has been a fellow of ASM for more than 30 years, and a member of the Russian Academy of Science, and was awarded the Ben Gurion Medal (Israel) and the Medal for Service to Powder Metallurgy by the Metal Powder Association. Recently, he was awarded an honorary doctorate in engineering by the University of Sheffield (United Kingdom) for his work in the field of titanium science and technology. Dr. Ma Qian is a Professor of Advanced Manufacturing and Materials in the Royal Melbourne Institute of Technology (RMIT University) in Melbourne, Australia, and an Honorary Professor of Materials Engineering in The University of Queensland in Brisbane, Australia. He received both his B.Eng. (1984) and D.Eng. (1991) in Foundry Metallurgy from Beijing University of Iron and Steel Technology (the present University of Science and Technology Beijing). His current research centers on metal additive manufacturing (AM), solidification, and powder metallurgy (Ti and Al alloys). With his collaborators and students, he has published 202 journal papers in these areas with nearly 5000 Scopus citations. He co-authored the fifth edition of

xxiv

About the editors

the Ian Polmear book, Light Alloys: Metallurgy of the Light Metals (2017, Elsevier), with Ian Polmear, David St. John, and Jianfeng Nie, and co-edited Titanium Powder Metallurgy: Science, Technology and Applications (2015, Elsevier) with F. H. Froes. He received the Australian CAST-CRC Industry Partner’s Award (2002, inaugural), the TMS Magnesium Technology Award (2003), the ASM Henry Howe Medal (2006), the Australian ARC CoE for Design in Light Metals best paper awards (2012, 2103), and five other awards. In particular, he co-developed “The Interdependence Theory for Grain Formation in Solidification” (Acta Mater. 2011;59:4907; 2010;58:3262), which is being applied to metal AM today, and a novel selective laser melting (SLM) process for Ti-6Al-4V (Acta Mater. 2015;85:74). He initiated the biannual international conference on Titanium Powder Metallurgy in 2011 (co-sponsored by Materials Australia, TiDA, TMS, JSPM, and CSM), which has been successively held in 2013 (Waikato, New Zealand), 2015 (L€uneburg, Germany), and 2017 (Xian, China), and co-chaired the first Asia-Pacific International Conference on Additive Manufacturing (December 4–6, 2017, RMIT University, Melbourne). Currently, he serves as an Associate Editor for Acta Materialia and Scripta Materialia.

Titanium for medical and dental applications—An introduction

1.1

F.H. (Sam) Froes Consultant to the Titanium Industry, Tacoma, WA, United States

1.1.1

Background

Titanium was first used in orthopedic applications in the 1950s. Orthopedics is that branch of surgery which is concerned with conditions involving the musculoskeletal system. Orthopedic surgeons use both surgical and nonsurgical means to treat mus culoskeletal trauma, spine diseases, sports injuries, and much more. Currently, titanium alloys are the main preference (Fig. 1.1.1) for orthopedic devices such as hip joints, bone screws, knee joints, spinal fusion cages, shoulder and elbow joints, and bone plates and scaffolds (Fig. 1.1.2) [1–3]. Titanium is the metal of choice for orthopedic practitioners because of its general corrosion resistance (Fig. 1.1.3) unlike iron (Fig. 1.1.4). It is inert in the human body and is resistant to attack by body fluids. Additionally, it has proven to be compatible with bone density, it is strong, and has a low modulus; hence, it is an excellent material for the orthopedic arena. Titanium is also used in cardiovascular applications (i.e., pertaining to the heart and blood vessels throughout the body). Blood vessels carry nutrients and oxygen to the tissues of the body and remove carbon dioxide and other wastes. They are subject to a number of diseases including arteriosclerosis, coronary artery disease, heart valve dis ease, and many more. The human body readily accepts titanium; it has proven to be more biocompatible than stainless steel or cobalt chrome. In addition, titanium has a higher fatigue strength than many other metals. It is also compatible with MRI (magnetic resonance imaging) and CT (computed technology), which also contributes to making it the material of choice in orthopedic applications. More than 1000 tonnes (2.2 million pounds) of titanium are implanted in patients worldwide every year, and this is set to increase as people live longer, or are seriously injured in road traffic or other accidents. Titanium is light, strong, and totally biocompatible, making it one of the few materials currently known that naturally matches the requirements for implantation in the human body. Because titanium resists corrosion, is biocompatible, and has an innate ability to join with human bone, it has become a staple of the medical field. From surgical titanium instruments to orthopedic titanium

Titanium in Medical and Dental Applications. https://doi.org/10.1016/B978-0-12-812456-7.00001-9 © 2018 Elsevier Inc. All rights reserved.

4

Fig. 1.1.1 Characteristics of titanium which make it attractive for medical and dental applications.

Titanium in Medical and Dental Applications

High corrosive resistance Low specific gravity

Biocompatible material

Nonmagnatic property

TITANIUM

High specific strength

Fig. 1.1.2 Ti-6Al-4V scaffold for medical implant applications fabricated using Selective Laser Melting Additive Manufacturing technology.

rods, pins, and plates, titanium has become the fundamental material used in medicine. Titanium is also much-used outside medicine. This makes it incredibly useful for many different industries, including the automotive, aerospace, and architectural worlds.

Titanium for medical and dental applications—An introduction

5

Fig. 1.1.3 General corrosion behavior of commercial purity titanium and Ti-Pd alloys compared to other metals and alloys in oxidizing and reducing acids; with and without chloride ions. In general, each metal or alloy can be used for those environments below its respective solid lines.

Fig. 1.1.4 Titanium is very corrosion resistant, unlike iron [4].

1.1.2

Body implants

The implant must be permanent in critical applications where it cannot readily be maintained or replaced. There is no more challenging situation in this respect than implants in the human body. Here, the effectiveness and reliability of both

6

Titanium in Medical and Dental Applications

implants and medical and surgical instruments are essential factors in saving lives and in the long-term relief of suffering and pain. Implantation represents a potential assault on the chemical, physiological, and mechanical structure of the human body. There is nothing comparable to a metallic implant in living tissue. Most metals in body fluids and tissue are found in stable organic complexes. Corrosion of implanted metal by body fluids results in the release of unwanted metallic ions, with likely interference in the processes of life. Corrosion resistance is not itself sufficient to suppress the body’s reaction to cell-toxic metals or allergenic elements such as nickel, and even in very small concentrations from a minimum level of corrosion, these may initiate rejection reactions. Titanium is judged to be completely inert and immune to corrosion by all body fluids and tissue, and is, thus, wholly biocompatible. The reasons for selecting titanium for implantation are a combination of the most favorable characteristics, including immunity to corrosion, biocompatibility, strength, low modulus and density, and the capacity for joining with bone and other tissue (osseointegration). The mechanical and physical properties of titanium alloys combine to provide implants which are highly damage tolerant. The human anatomy naturally limits the shape and allowable volume of implants. The lower modulus of titanium alloys compared to steel is a positive factor in reducing bone resorption. Two further parameters define the usefulness of the implantable alloy, the notch sensitivity (i.e., the ratio of tensile strength in the notched to the unnotched condition) and the resistance to crack propagation, or fracture toughness. Titanium scores well in both cases. Typical NS/TS ratios for titanium and its alloys are 1.4–1.7 (1.1 is a minimum for an acceptable implant material). The fracture toughness of all high-strength implantable alloys is above 50 MPa.m-½, with critical crack lengths well above the minimum for detection by standard methods of nondestructive testing. Examples of titanium use in body implants are shown in Figs. 1.1.5–1.1.7, and its dental use is illustrated later in this chapter. For dental applications, a major change in restorative dental practice worldwide has been made possible using titanium implants. A titanium “root” is introduced into the jaw bone with time subsequently allowed for osseointegration. The superstructure of the tooth is then built onto the implant to give an effective replacement. Titanium is also used in surgery to repair facial damage. Use of the patient’s own tissue cannot always obtain the desired results. Artificial parts may be required to replace facial features lost through damage or disease and so restore the ability to speak or eat, as well as improve cosmetic appearance. Osseointegrated titanium implants meeting all the requirements of biocompatibility and strength have enabled unprecedented advances in surgery, for the successful treatment of patients with large defects and hitherto highly problematic conditions. A comprehensive article on “Understanding Implants in Knee and Hip Replacement” appeared in the journal of the International Titanium Association 2nd Quarter, 2016 Medical Edition [5]. It showed that titanium knee replacements can be tailored to the patient’s size, weight, and gender (a woman has a narrower bone structure, especially on the femur). A companion article [5, p. 20] discusses the

Titanium for medical and dental applications—An introduction

7

Fig. 1.1.5 Knee joint implant replacement x-ray showing in medical orthpodedic traumatology scan. Courtesy of Shutterstock.

Fig. 1.1.6 Titanium bone implants.

influence of processing (in particular, etched versus unetched conditions) on the performance of titanium middle ear prostheses, with the etched condition exhibiting superior performance. A further article [5, p. 76] presents details on a 3D-printed titanium hip implant with a fully porous cup (Fig. 1.1.8) allowing in-growth of bone and tissue, resulting in superior performance. An additive manufactured rib cage and sternum implant is shown in Fig. 1.1.9A and B [5, p. 82]. Fig. 1.1.10 illustrates a fecal continence restoration system consisting of a series of titanium beads with magnetic cores connected by titanium wires to form a ring [5, p. 83].

8

Titanium in Medical and Dental Applications

Fig. 1.1.7 A traditional total hip replacement implant. From http://www.geripal.org/2013/01/metal-on-metalhip-replacements-tragic.html.

Fig. 1.1.8 Smith and Nephew revised acetabular fully porous cup with Conceloc technology. (Courtesy Professor H. P. Tang, State Key Laboratory of Porous Metal Materials, Northwest Institute for Nonferrous Metal Research).

An additive manufactured jaw bone is shown in Fig. 1.1.11 [4]. Other interesting work involved a study of Ti-Au alloys [6]. An alloy consisting of 3:1 Ti::Au (Fig. 1.1.12) was found to be four times harder than pure titanium and was more biocompatible and exhibited superior wear resistance, making it appear to be an attractive choice for implant applications.

Fig. 1.1.9 (A) A 3D additive manufactured rib cage and sternum implant. (B) How the component shown in A is inserted into the human body.

Fig. 1.1.10 Titanium fecal continence restoration system.

10

Titanium in Medical and Dental Applications

Fig. 1.1.11 An additive manufactured jaw bone.

Fig. 1.1.12 Crystal structure of beta titanium-3 gold (Courtesy E. Morosan, Rice University).

1.1.3

Dental implants

The use of titanium in dental applications has also increased dramatically over the past 20 years. The replacement of missing teeth with implant-supported prostheses, illustrated in Figs. 1.1.13–1.1.14 [4] has become widely accepted in dentistry for the rehabilitation of fully and partially edentulous patients. This breakthrough in oral rehabilitation is based on the concept of osseointegration. This biological phenomenon is described as direct bone deposition upon a titanium implant surface. Currently, commercially pure (CP) titanium has become the material of choice in implant dentistry, since it has excellent biological and biomechanical properties.

Titanium for medical and dental applications—An introduction

11

Fig. 1.1.13 Titanium dental implants (top) and their insertion into the mouth (bottom).

Titanium implant description Typical ‘two-piece’ titanium implant connection that is ready to have the crown or bridge connected with a little screw.

Titanium implants are placed deep under the gum in order to hide the metal color. This is difficult to brush and for this reason, the connections tend to produce halitosis. This is the area where the corrosion of metal implants is greater. This connections have been proven to accumulate bacteria in all kinds of situations with poor or very good oral hygiene and in any kind or brand of titanium two-piece implants. This is the location where the crown or bridge is connected to the titanium implants

Fig. 1.1.14 How the dental implant procedure works.

1.1.4

Titanium surgical instruments

A wide range of surgical instruments are made from titanium (see Figs. 1.1.15–1.1.17 [4]). One of its advantages in this respect is its lightness, which helps to reduce the surgeon’s fatigue. Instruments are frequently anodized to provide a nonreflecting

12

Titanium in Medical and Dental Applications

Fig. 1.1.15 Titanium surgical instruments, scissors, forceps, and needle holders.

Fig. 1.1.16 Titanium instruments, surgical screws, and various implants.

surface, essential in microsurgical operations such as eye surgery. Titanium instruments can be sterilized repeatedly without compromising edge or surface quality, corrosion resistance, or strength. Titanium is nonmagnetic, and there is, therefore, no threat of damage to small, sensitive, implanted electronic devices.

Titanium for medical and dental applications—An introduction

13

Fig. 1.1.17 Additional titanium instruments.

1.1.5

Titanium in wheel chairs, etc.

The same characteristics that make titanium a preferred choice for implants and instruments make it a good choice, particularly in tubular form (Fig. 1.1.18A and B [4]) for wheel chairs (Fig. 1.1.19 [4]) and walkers (Fig. 1.1.20 [4]). The tubular products are generally fabricated as either seamless or welded components from commercially pure titanium or the Ti-3Al-2.5 V alloy.

1.1.6

Specifications for titanium in medical and dental applications

Forms and material specifications are detailed in a number of international and domestic specifications, including ASTM and BS7252/ISO 5832, as shown in Table 1.1.1 [2]. Alloys such as Ti-6Al-7Nb eliminate the use of V, an element which can cause cytotoxic outcomes (negatively influencing human cells in a similar manner to the effect of puff adder venom). Mechanical properties of titanium alloys suitable for use in medical and dental applications are shown in Table 1.1.2 [7].

1.1.7

Other titanium-based materials

Nickel titanium, also known as nitinol (part of a shape memory alloy), is a metal alloy of nickel and titanium, in which the two elements are present in roughly equal atomic percentages, for example, nitinol 55, nitinol 60. Nitinol alloys exhibit two closely related and unique properties: shape memory effect (SME) and superelasticity (SE; also called pseudoelasticity, PE). Shape memory is the ability of nitinol to undergo deformation at one temperature, then recover its original, undeformed shape upon heating above its “transformation temperature.” Superelasticity occurs at a narrow

14

Titanium in Medical and Dental Applications

Fig. 1.1.18 (A) Titanium tubing. (B) Titanium tubing ready for use in wheel chair or walker applications.

Titanium for medical and dental applications—An introduction

Fig. 1.1.19 (A,B) Titanium is a popular choice for wheel chairs.

Fig. 1.1.20 A titanium walker.

15

16

Titanium in Medical and Dental Applications

Table 1.1.1 Specifications for titanium in medical and dental applications ASTM

BS/ISO

Alloy(S) designation

F67 F136 F1472

Part 2 Part 3 Part 3

F1295 – F1580 F1713 F1813

Part 11 Part 10 – – –

Unalloyed Titanium—CP grades 1–4 (ASTM F1341 specifies wire) Ti6Al4V ELI wrought (ASTM F620 specifies ELI forgings) Ti6Al4V standard grade (SG) wrought (F1108 specifies SG castings) Ti6Al7Nb wrought Ti5Al2.5Fe wrought CP and Ti6Al4V SG powders for coating implants Ti13Nb13Zr wrought Ti12Mo6Zr2Fe wrought

Table 1.1.2

Titanium alloys suitable for medical applications ASTM grade

Property

1

2

3

4

5

Yield strength (MPa) Ultimate tensile strength (MPa) Elongation (%) Elastic modulus (GPa)

170 240

275 345

380 450

483 550

795 860

24 103–107

20 103–107

18 103–107

15 103–107

10 114–120

Adapted from ASTM F67 (Grade 1 to 4) and F136 (Grade 5).

temperature range just above its transformation temperature; in this case, no heating is necessary to cause the undeformed shape to recover, and the material exhibits enormous elasticity, some 10–30 times that of ordinary metal. Nitinol is highly biocompatible and has other properties that make it suitable for use in orthopedic implants. Its unique properties have caused high demand in less invasive medical devices. Nitinol tubing is commonly used in catheters, stents, and superelastic needles. In colorectal surgery, the material is used in devices for reconnecting the intestine after removing the pathology. Nitinol is used for devices developed by Franz Freudenthal to treat patent ductus arteriosus, blocking a blood vessel that bypasses the lungs and has failed to close after birth in an infant. In dentistry, this alloy is used in orthodontics for brackets and wires connecting the teeth. Once the SME wire is placed in the mouth, its temperature rises to ambient body temperature. This causes the nitinol to contract back to its original shape, applying a constant force to move the teeth. These SME wires do not need to be retightened as often as other wires because they can contract as the teeth move, unlike conventional, stainless-steel wires. Additionally, nitinol can be used in endodontics, where nitinol files are used to clean and shape the root canals during the root canal procedure.

Titanium for medical and dental applications—An introduction

17

Nitinol’s unusual properties are derived from a reversible solid-state phase transformation known as a martensitic transformation. At high temperatures, nitinol assumes an interpenetrating, simple, cubic structure referred to as austenite. At low temperatures, nitinol spontaneously transforms to a more complicated, body-centered, tetragonal crystal structure known as martensite. The temperature at which austenite transforms to martensite is generally referred to as the transformation temperature. The mechanism of the shape memory effect involving the transformation from the low-temperature martensite to the higher-temperature austenite causes the original shape to be recovered (see Figs. 1.1.21 and 1.1.22 [4]).

15

nm

0 0.3

46

22

20

nm

0.3015 nm

0.

96.8

41

ᑻ

0.

Austenite

nm

Martensite

Fig. 1.1.21 Mechanism of the shape memory effect. The transformation from the lowtemperature martensite to the higher-temperature austenite causes the original shape to be recovered.

Fig. 1.1.22 Demonstration of the shape memory effect.

18

Titanium in Medical and Dental Applications

Fig. 1.1.23 A nitinol stent.

Fig. 1.1.24 How a nitinol stent works.

A nitinol stent and how it works is shown in Figs. 1.1.23 and 1.1.24 [4]. A nitinol fixation device is shown in Fig. 1.1.25 [4]. Dental use of nitinol is shown in Fig. 1.1.26, [4] a dental brace. Recently, alloys have been developed beyond the first-generation terminal alloys. These include Ti-6Al-4V, Ti-6Al-7Nb, and the commercially pure grades. These alloys are low modulus near beta and beta alloys, which match the low modulus of human bone much better [8]. Alloys such as Ti-29Nb-13Ta-4.6Zr, Ti-13Nb-13Zr, and Ti-10Nb-10Zr have been fabricated by metal-injection molding methods and have good strengths and excellent biocompatibility with the human body.

Titanium for medical and dental applications—An introduction

19

Fig. 1.1.25 Nitinol fixation devices.

Fig. 1.1.26 Nitinol dental braces. (A) Schematic of braces in place on teeth and (B) dental braces on teeth.

1.1.8

Post script

A comprehensive summary of the use of AM in dentistry is shown in Fig. 1.1.27.

1.1.9

This book

This book focuses on titanium in medical and dental applications, including implants, instruments, and devices such as wheel chairs and walkers. The book contains information on various fabrication techniques including conventional ingots and

20

Titanium in Medical and Dental Applications

3D Printing—Manufacturing Methods—Dental Industry

Laminated Object Manufacturing (LOM) 3.2%

n itio ) os ep FDM dD ( se ling Fu ode 0.8% 1 M

Sele Me ctive L ltin ase g 11. (SLM r ) 3%

3D in

kjet P 7.1% rinting

rs he Ot .4% 8

Se tive Laser lec Sintering (SL S) 14.3%

La Dire se ct rS M 2. int eta 3 % er l in g

A) (SL phy gra o h lit reo 27.5% Ste

) ht ig LP l L (D ta ing i ig D ess .5% oc 1 Pr c Beam Electroni BM) (E Melting 2% 3.

Others Include Laser Powder Sintering | Solid Freeform Fabrication | Robocasting

Fig. 1.1.27 A comprehensive summary of the use of AM in dentistry. Courtesy of Sagacious Research.

subsequent processing (forging, rolling, extrusion, etc.), and approaches such as powder metallurgy, additive manufacturing, and metal injection molding. It also contains information on a number of alloys starting with commercially pure titanium and Ti-6Al-4V and extends to compositions such as nitinol (with compositions close to equiatomic Ti-Ni).

References [1] [2] [3] [4]

Titanium Industries, Inc., US office of Technical Assessment Web Page (accessed 7-07-16). AZO Materials Web Page (accessed 7-07-1) Wikipedia, (accessed 7-17-16). Internet Explorer, Titanium Implants Illustrations, vol. 8 (accessed 11-25-16).

Titanium for medical and dental applications—An introduction

21

[5] Padgett, D.E., Windsor, R.E., 2016. International Titanium Association. Titan. Today (2nd Quarter), 11. [6] Anon, 2016. Gold boosts titanium knee strength. Adv. Mater. Process. 174 (8) 8,9. [7] Elias, C.N., et al., 2008. Biomedical applications of titanium and its alloys. JOM (March), 46. [8] B. Williams, Powder injection molding international, “World PM2016: PIM Technical Sessions Review Advances in Novel Titanium Alloys for Biomedical Applications.

Titanium background, alloying behavior and advanced fabrication techniques—An overview

1.2

F.H. (Sam) Froes*, M. Qian† *Consultant to the Titanium Industry, Tacoma, WA, United States, †RMIT University, Melbourne, VIC, Australia

1.2.1

Titanium alloys and their importance

Titanium alloys are among some of the most important new materials that are key to improved performance in aerospace and terrestrial systems (Figs. 1.2.1 and 1.2.2). They are even finding applications in the cost-aware automotive industry (Table 1.2.1) [2–10]. These applications are due to the excellent combinations of specific mechanical properties (properties normalized by density) and outstanding corrosion behavior of the titanium alloys. [2–10] However, widespread use is prevented by the high cost of titanium alloys compared to competing materials (Table 1.2.2). In this chapter, the behavior of titanium will be followed by a brief overview of the metallurgy of the titanium system and a discussion of two developing net or near-net shape techniques: Additive Manufacturing (AM) and Metal Injection Molding (MIM). The high cost of titanium compared with the other metals shown in Table 1.2.2 has resulted in the yearly consumptions shown in Table 1.2.3. Recent publications [2–10] show the cost of fabricating various titanium precursors and mill products (very recently, the price of TiO2 has risen to $2.00 per pound and TiCl4 to $0.55 per pound), and it has been pointed out that the cost of extraction is a small fraction of the total cost of a component fabricated by the cast and wrought (ingot metallurgy) approach. To make a final component, the mill products must be machined, often with very high buy-to-fly ratios (which can reach as high as 40:1). The generally accepted cost of machining a component is that it doubles the cost (with the buy-to-fly ratio being another multiplier in cost per pound). The breakdown of manufacturing costs in the case of the Boeing 787 side-of-body chord. Thus, while improvements in the machining of titanium have occurred anything that can be done to produce a component which is closer to the final configuration will result in a cost reduction – hence, the attraction of near-net shape powder metallurgy techniques [11–17] such as MIM [18–32] and AM [33–37].

Titanium in Medical and Dental Applications. https://doi.org/10.1016/B978-0-12-812456-7.00002-0 © 2018 Elsevier Inc. All rights reserved.

24

Titanium in Medical and Dental Applications

Fig. 1.2.1 The Guggenheim Museum in Bilbao, northern Spain, which is sheathed with titanium sheet. http://www.bbc.com/culture/ story/20131218-when-themuseum-is-the-art.

(A)

(B)

Natural tooth

Jawbone

(C)

Custom made crown

Titanium implant

(D)

Fig. 1.2.2 Titanium is used for a wide variety of items, such as bike (A) frames, (B) eyeglass frames, (C) watches and (D) dental implants. (A) https://delleks.en.taiwantrade.com/product/titanium-bike-frame-302130.html; (B) http:// static.zennioptical.com/images/product/52/57/525721_lg.jpg; (C) http://www.ebay.com/gds/ Top-5-Titanium-Watches-for-Men-/10000000177748096/g.html; (D) I. Polmear, D.H. St. John, J.F. Nie, M. Qian, Light Alloys: Metallurgy of Light Metals, fifth ed., Elsevier, 2017.

Titanium background, alloying behavior and advanced fabrication techniques—An overview

25

Some of the components in automobiles which have/are produced from titanium alloys

Table 1.2.1

Component

Material

Manufacturer

Model

Connecting rods

Ti-3Al-2V-rare earth Ti-6A1-4V Ti-6A1-4V

Honda

Acura NSX

Ferrari Porsche

Ti grade 2 Ti grade 1s Ti grade 1 Ti-6A1-4V Ti-6A1-4V & PM-Ti Ti-6AI-4V TIMETAL LCB Ti-6A1-4V β-titanium alloys γ-TiAl Ti grade 2 Ti-6AI-4V

Daimler Volkswagen Honda Porsche Toyota

All 12-cyl. Sport wheel option S-Class All S2000 Roadster GT3 Altezza 6-cyl.

Connecting rods Wheel rim screws Brake pad guide pins Brake sealing washers Gearshift knob Connecting rods Valves Turbo charger wheel Suspension springs Wheel rim screws Valve spring retainers Turbo charger wheel Exhaust system Wheel rim screws Valves Suspension springs

Table 1.2.2

Ti-6A1-4V & PM-Ti TIMETAL LCB

a

Nissan

Truck diesel Lupo FSl M-Techn. option All 1.8 1 – 4-cyl. Lancer Corvette Z06 Sport package GTI Infiniti Q45

Ferrari

360 Stradale

Cost of titanium—A comparisona

Item

Ore Metal Ingot Sheet

Daimler Volkswagen BMW Mitsubishi Mitsubishi General Motors Volkswagen

Material ($/pound) Steel

Aluminum

Titanium

0.02 0.10 0.15 0.30–0.60

0.01 1.10 1.15 1.00–5.00

0.22 (rutile) 5.44 9.07 15.00–50.00

2015 Contract prices. The high cost of titanium compared to aluminum and steel is a result of (a) high extraction costs and (b) high processing costs. The latter relates to the relatively low processing temperatures used for titanium and the conditioning (surface regions contaminated at the processing temperatures, and surface cracks, both of which must be removed) required prior to further fabrication).

26

Titanium in Medical and Dental Applications

Table 1.2.3

Metal consumption

Structural metals

Consumption/year (103 metric tons)

Ti Steel Stainless steel Al

50 700,000 13,000 25,000

1.2.2

Metallurgy of the titanium system

The chemistry (Fig. 1.2.3) [2–10] and microstructure (Fig. 1.2.4) [2–10] of the titanium alloy determine its chemical and physical properties. Alpha alloys are characterized by relatively low strength (
80 ksi UTS); a number of members of this class of alloys are used for high-temperature applications (up to 600°C). Equiaxed alpha and

Fig. 1.2.3 Compositions of US technical alloys mapped onto a pseudobinary β-isomorphous phase diagram. Courtesy ASM Int.

Titanium background, alloying behavior and advanced fabrication techniques—An overview

27

Fig. 1.2.4 Microstructure of Ti-6Al-2Sn-4Zr-2Mo: (A) β worked followed by α-β anneal to produce lenticular α morphology; (B) α-β worked and α-β annealed to give a predominantly equiaxed α shape; (C) α-β worked followed by duplex anneal: just below the β transus temperature (reduced volume fraction of equiaxed α compared to (B)), and significantly below the β transus temperature (to form the lenticulary α between equiaxed regions.

elongated alpha are the two basic microstructures exhibited by conventional alpha titanium alloys. They exhibit good ductile behavior and fatigue crack initiation on the one hand and good fracture toughness and creep performance on the other. The alpha-beta class of alloys has higher strength in combination with reasonable levels of ductility. For example, the Ti-6Al-4V alloy (the work-horse alloy of the titanium industry) exhibits minimums of 130 ksi UTS and 12% elongation. The beta alloys have strength equivalent to or higher than that of Ti-6Al-4V with significantly higher ductility. A class of titanium alloys not shown on Fig. 1.2.3 are the intermetallic TixAl (x ¼ 1 or 3) alloys, which have excellent high-temperature behavior, but very low room-temperature ductility (often 2% elongation maximum). Generally, titanium alloys increase in strength and decrease in ductility as the oxygen level is increased [2–10]. The maximum aerospace specification for oxygen Ti-6Al-4V is 0.20 wt.%. Research has shown that powder metallurgy Ti-6Al-4V for non-aerospace applications can retain good tensile ductility when the oxygen content is below about 0.32 wt.%. [10]

1.2.3

Advanced fabrication techniques for titanium components

Traditionally, titanium parts are typically fabricated by subtractive techniques, i.e., by machining of mill products (Fig. 1.2.5). However, net or near-net methods of component production, such as MIM and AM, significantly reduce the amount of machining required resulting in lower cost high integrity products.

1.2.3.1 Metal injection molding of components Metal injection molding of components is particularly applicable to complex small parts (generally weighing 1 pound or less) of high complexity (Fig. 1.2.6), [18–32] the smaller “chunky” automobile parts listed in Table 1.2.1 are components

28

Titanium in Medical and Dental Applications

Fig. 1.2.6 Diagram showing where Ti MIM is most appropriately used in comparison with other fabrication processes. From Miura H., Osada T., Itoh Y. (2015) Metal Injection Molding (MIM) Processing. In: Niinomi M., Narushima T., Nakai M. (Eds.) Advances in Metallic Biomaterials. Springer Series in Biomaterials Science and Engineering, vol 4. Springer, Berlin, Heidelberg.

Production run (thousands)

Fig. 1.2.5 A titanium component fabricated by 5-axis machining.

Die casting Fine blanking Forging P/M

Optimum region for PIM

Investment casting Machining Low

Medium Part complexity

High

which could potentially be produced by the MIM technique. Fig. 1.2.7 shows the MIM market in Japan for 2015. MIM Ti accounted for 2.8% of this market. Fig. 1.2.8 illustrates the basic steps involved in a typical MIM process [20]. First, a polymer binder and metal powder are mixed to form the feedstock, which is molded, debonded, and sintered. The process relies on the thermoplastic binder for shaping at a moderately elevated temperature of about 150°C. Fig. 1.2.9 shows several MIM Ti parts fabricated from non-spherical hydride-dehydride (HDH) Ti powder. MIM works best in situations where there are a large number of small, complex parts. This can be further classified where Ti MIM includes a mixture of cosmetic parts (where the mechanical properties are not important) and structural parts (exposed to stress, making mechanical properties of importance). A major contributor to the mechanical properties of commercially pure titanium (CP Ti) is the interstitial levels, particularly oxygen. Thus, the aerospace oxygen specification for CP Ti (Grade 4) is 0.4 w/o, whereas for Ti-6A l-4V it is 0.2 w/o; with the former composition being used at lower strength (Grade 4) (
80 ksi) and the latter at levels of 130–140 ksi. Thus, “cosmetic” parts such as watch cases are fabricated from CP Ti.

Titanium background, alloying behavior and advanced fabrication techniques—An overview

Fig. 1.2.7 MIM markets in Japan for 2015. Reproduced with permission from Japan Powder Metallurgy Association.

29

30

Titanium in Medical and Dental Applications

Fig. 1.2.8 Schematic of the steps involved in powder injection molding (see text for details) [26].

Fig. 1.2.9 Titanium parts fabricated by metal injection molding from non-spherical hydride-dehydride (HDH) titanium powder. The dimensions of the parts are 21 mm  17 mm  8 mm. Courtesy Guangzhou Xu Peng Technology Co., Ltd.

The metal powder injection molding process was developed from the injection molding of plastics, a process developed for long production runs of small (normally below 400 g), complex-shaped parts in a cost-effective manner. By increasing the metal (or ceramic) particle content, the process developed into one for production of high-density metal, intermetallic, or ceramic components (Fig. 1.2.8). [26] Early attempts to produce a viable titanium MIM process were plagued by the lack of suitable powder, as well as inadequate protection of the titanium during elevatedtemperature processing and less-than-optimum binders for a material as reactive as titanium. [2–10] However, some MIM practitioners have now learned what the titanium community has long known—that titanium is the universal solvent and must be treated accordingly. Suitable powders are now available and sintering furnaces viable for use with titanium are now in place. Thus, the challenge now is to find suitable binders. Unfortunately, some polymer binders which are known for their ability to readily thermally unzip to their starting monomers impurities into the sintered Ti MIM

Titanium background, alloying behavior and advanced fabrication techniques—An overview

31

bodies are not suitable because their depolymerization occurs close to the temperatures where impurity uptake initiates (at  260°C). Alternative binder systems, based on catalytic decomposition of the polyacetals, are promising but require expensive capital-intensive equipment to handle the acid vapor catalyst. The process also requires a suitable means of eliminating the formaldehyde oligomer that forms as a polymer decomposition byproduct. However, there are a number of binder systems which appear to have the necessary characteristics to be compatible with titanium (Table 1.2.4) giving acceptable levels of oxygen content, strength and ductility (Table 1.2.5). Factors which effect the strength—ductility are the oxygen level

Binder systems which appear to be compatible with Ti-6Al-4Va

Table 1.2.4

Polypropylene-ethel vinyl acetate-paraffin wax-carnauba wax-diocytl phthalate [22] Polyethylene, paraffin and stearic acid [23] Polypropylene-polymethyl methacrylate-paraffin-stearic acid [24] Polypropylene-paraffin-carnauba, etc. [25] Secret [26] Naphthalene-stearic acid-ethylene vinyl acetate [27] Paraffin wax-polyethylglycol-polyethylene-stearic acid [28] Paraffin wax-copolymer-stearic acid [29] Atactic polypropolene-carnauba wax-paraffin wax-stearic acid [30] PP-EVA-PW-CW-DOP [31] Specially developed on polymer base [32] a

See Table 1.2.5.

Table 1.2.5

Characteristics of Ti MIMa

Oxygen content (w/o)

Relative density (%)

UTS (ksi)

Elongation (%)

Reference

0.24 0.28 0.25–0.28 0.20 0.19 0.17 0.54 – 0.32 0.34 0.24

96.0 97.1 95.5 95.1 95% 99.5d 96.7 99.5 >96.0 96 98

140.8 118.3 121.9 94.3 – 152.1 121.9 136.1 136.4 139.3 133.5

12 7.8b 14.0 22c – 14.6 9.0 14.0 2.5 11.2 14.0e

[22] [23] [24] [25] [26] [27] [28] [29] [30] [31] [32]

a

Ti-6Al-4V unless stated. Ti-6Al-7Nb. Commercially pure titanium. d HIP’d. e Near alpha alloy. b c

32

Titanium in Medical and Dental Applications

(strength up, ductility down), relative density (strength and ductility up), and the beta grain size (smaller grain size for increased strength and ductility). An additional factor is powder size, with a smaller size likely to give increased oxygen content and a finer beta grain size. Currently, titanium MIM parts are made up to a foot in length, but parts over three or four inches (about 50 g in weight) are not common. The limiting factors at this time are dimensional reproducibility and chemistry. Due to shrinkage, large parts become dimensionally more difficult to make due to loss of shape. If the parts have flat surfaces which rest on the setter, they come out fairly consistently. In contrast, parts with multiple surfaces that require setters in complex shapes become less reproducible as the size goes up. Further, large overhanging areas become difficult to control dimensionally due to gravity. With increasing experience, the packing density of titanium powder mixes will increase, especially as new binders become available, meaning that shrinkage can decrease making the dimensional problems less of a factor.

1.2.3.2 Additive manufacturing The cost of fabricating various titanium precursors and mill products has been discussed in publications over the past few years [33–37] (very recently the price of TiO2 has risen to $2.00 per pound and TiCl4 to $0.55 per pound), and it has been pointed out that the cost of extraction is a small fraction of the total cost of a component fabricated by the cast and wrought (ingot metallurgy) approach. To reach a final component, the mill products must be machined, often with very high buy-to-fly ratios (which can reach as high as 40:1). The generally accepted cost of machining a component is that it doubles the cost. A wide variety of metal AM processes have been developed to date (Fig. 1.2.10) [35]. The two basic approaches are Powder Bed Fusion (PBF) and Direct Energy Deposition (DED). The PBF technique permits the fabrication of complex features, hollow cooling passages, high precision parts, and single metal builds. The DED approach allows large build envelopes, high deposition rates, multiple materials, and addition of material to existing components. Mechanical properties are at least at Ingot Metallurgy levels (including fracture toughness) as shown in Fig. 1.2.11 [18–32]. A wide variety of titanium medical components have been fabricated by AM. Fig. 1.2.12 show three examples of this technique.

1.2.3.3 The future of MIM and AM The markets for both MIM Ti and AM Ti techniques are likely to grow. Several factors which will impact that growth will be availability of low-cost (less than $20/lb or $44/kg) powder of the appropriate size (less than 40 microns) and good purity (which maintained throughout the fabrication process). For non-aerospace applications, the purity level of the Ti-6Al-4V alloy can be less stringent. For instance; e.g., the oxygen level can be up to 0.3 wt% while still exhibiting acceptable ductility levels while within aerospace, there is a requirement of a maximum oxygen level of

Metal additive manufacturing processes

Direct metal deposition (fusion)

Powder bed fusion, sintering or binding

Wire/filament deposition

Metal powder deposition

Electron beam

Laser

Selective melting SLM Solutions ReaLizer 3D Systems (former Phenix systems) Renishaw Concept Laser EOS (DMLS) TRUMPF Matsuura (hybrid) Additive Industries (MetalFAB1)

Selective sintering

Selective melting

Binder jetting, and sintering

With binder

3D systems

Arcam

Plasma beam

Binder

ExOne Desktop metal

OPTOMEC - Laser engineered net shaping (LENS) BeAM - Laser metal deposition TRUMPF - TruLaser deposition DM3D (POM group inc.) - Direct Metal Deposition (DMD) IREPA LASER Laser CLAD Accufusion - Laser consolidation

Electron beam

Laser

Sciaky

Fig. 1.2.10 Various metal additive manufacturing processes developed to date [35].

TRUMPF

Liquid metal printing

Cold spray deposition

Metal foils or sheets

Metal nanoparticles suspended in a proprietary liquid

Metal powder

TIGtorch Pressure assisted solid state ultrasonic welding

Wire deposition

Norsk titanium

Ultrasonic AM (non fusion)

Nuclear AMRC

Soniclayer

Nano Metal Jetting©

Xjet

High speed particles in a hot gas stream

DYMET ASB MEC KINETIKS Dentaco

Titanium in Medical and Dental Applications

YS

Elongation 30

1200

25

1000

20

800

15

600

10

400

Wrought, annealed

Forged

Cast

EBM EBM, HIP

DMLS

DMLS, HIP+HT

0 LENS, HT

0 LENS, HIP

5 DMD

200 DMD, HIP+HT

Strength (MPa)

UTS 1400

Elongation (%)

34

Fig. 1.2.11 Tensile strength, yield strength, and elongation of Ti-6Al-4V alloy built using various AM processes. DMD, direct metal deposition; LENS, laser engineered net shaping; DMLS, direct metal laser sintering; EBM, electron beam melting; HIP, hot isostatic pressing; HT, heat treatment.

Fig. 1.2.12 Ti-6Al-4V cages for medical implant applications, fabricated using Selective Laser Melting technology, that have received FDA approval. EIT Cellular Titanium.

0.2 wt%) [2–10]. For commercially pure (CP) grades, oxygen levels can be even higher; up to at least 0.4 wt. % (Grade 4 CP titanium has a specification limit of 0.4 wt%.) [2–10] In fact, Grade 4 CP titanium (UTS 550 MPa [8 ksi]) while having a lower strength than regular Ti-6Al-4V (UTS 930 MPa [135 ksi]) may well be a better choice for the many potential MIM parts where cost is important. The Grade 4 would allow use of a lower-cost starting stock and a higher oxygen content in the final part. In addition, non-spherical low-cost hydride-dehydride (HDH) Ti powder or non-spherical TiH2 powder could be used for MIM as well (Fig. 1.2.8) [26] Further into the future, the beta alloys with their inherent good ductility (bcc structure) and the intermetallics with attractive elevated temperature capability are potential candidates for fabrication via MIM and AM. The science, technology, and cost

Titanium background, alloying behavior and advanced fabrication techniques—An overview

35

now seem to be in place for the titanium marketplace for both approaches to show significant growth. There have also been a number of developments, including development of suitable binders and sintering furnaces for MIM, which should lead to a reasonable growth of titanium products produced by this method. The biggest growth potential is in small complex shaped parts using the MIM approach. With the production of high integrity (particularly oxygen within specification limits) cost effective, complex MIM components a market in both aerospace and terrestrial industries should grow quite rapidly.

1.2.4

Conclusions

The significance of titanium in a number of market places (medical and dental, aerospace, automotive, and consumer products) has been reviewed. This was followed by a brief discussion of the behavior of titanium alloys (including phase diagrams) and a presentation of two advanced fabrication techniques for titanium components: Metal Injection Molding and Additive Manufacturing. Ways of producing titanium MIM parts have been discussed. Parts produced have historically been high in oxygen preventing their use in structural (load-bearing) applications; however, cosmetic parts (not exposed to any stress) have been successfully produced mainly from commercially pure grades of titanium (which allow up to 0.4 wt. %). Starting powders of suitable quality and price, along with sintering furnaces which minimize oxygen pick up, and a number of binders which do not result in significant oxygen pick up, are now available. Additive manufactured parts are currently being fabricated for various applications because of their short lead time, cost-effectiveness (resulting from the near-net shape capabilities), and acceptable mechanical properties. Thus, there is a pathway for increased production of structural titanium MIM and AM medical parts as well as surgical instruments and dental parts, and parts for the aerospace and automotive industries.

Acknowledgments The authors recognize useful discussions with Eric Bono, Bhaskar Dutta, Serge Grenier, Joe Grohowski, Andy Hanson, John (Qiang) Li, Tim McCabe, Eric Nyberg, Kevin Simmons and Fred Yolton. We also acknowledge the contribution of Ms. Marlane Martenick in helping formatting and typing the text.

References [1] P. Kitten, Bypass Fan Module for Engine Nacelles—Adds Bypass Thrust, Allows Stock Turbofan Engines, http://forum.kerbalspaceprogram.com/index.php?/topic/122139bypass-fan-module-for-engine-nacelles-adds-bypass-thrust-allows-stock-turbofan-engines/. (Accessed 23 July 2017).

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Titanium in Medical and Dental Applications

[2] Congress of the U.S. Office of Technology Assessment, Advanced Materials by Design, June 1988. [3] Materials Science and Engineering—Forging Stronger Links to Users, Publication NMAB-492, NMAB, National Academy Press, Washington, DC, 1999. [4] F.H. Froes, D. Eylon, H. Bomberger (Eds.), Titanium Technology: Present Status and Future Trends, TDA, Dayton, OH, 1985. [5] F.H. (Sam) Froes, T.-L. Yau, H.G. Weidenger, Titanium, zirconium and hafnium, in: K.H. Mataucha (Ed.), Materials Science and Technology—Structure and Properties of Nonferrous Alloys, VCH Weinheim, FRG, 1996, p. 401 (Chapter 8). [6] F.H. (Sam) Froes, Titanium, in: P. Bridenbaugh (Ed.), Encyclopedia of Materials Science and Engineering, Elsevier, Oxford, 2001, pp. 9361–9374 (subject editor). (Chapters 3.35a–3.3.5e). [7] F.H. (Sam) Froes, Titanium alloys, in: J.K. Wessel (Ed.), Handbook of Advanced Materials, Wiley Interscience, 2004, p. 271 (Ed. In Chief ). (Chapter 8). [8] R.R. Boyer, G. Welsch, E.W. Collings (Eds.), Materials Properties Handbook: Titanium Alloys, ASM Int., Materials Park, OH, 1994. [9] F.H. (Sam) Froes (Ed.), Titanium Physical Metallurgy, Processing and Applications, ASM, Materials Park, OH, February 2015. [10] M. Yan, P. Yu, G.B. Schaffer, M. Qian, Secondary phases and interfaces in a nitrogenatmosphere sintered Al alloy: TEM evidence for the formation of AlN during liquid phase sintering, Acta Mater. 58 (2010) 5667–5674. [11] F.H. Froes, D. Eylon, Powder metallurgy of titanium alloys, Int. Mater. Rev. 35 (1990) 162. [12] F.H. Froes, C. Suryanarayana, Powder processing of titanium alloys, in: A. Bose, R. M. German, A. Lawley (Eds.), Reviews in Particulate Materials, Vol. I, MPIF, Princeton, NJ, 1993, p. 223. [13] F.H. Froes, Introduction: developments to date, in: M. Qian, F.H. Froes (Eds.), Titanium Powder Metallurgy: Science, Technology and Applications, Butterworth-Heinemann, Elsevier, 2015. [14] F.H. Froes, Titanium powder metallurgy: developments and opportunities in a sector poised for growth, Powder Metall. Rev. 2 (4) (2013) 29–43. Winter. [15] F.H. (Sam) Froes, Titanium powder metallurgy: a review – part 1, Adv. Mater. Process. 170 (9) (2012) 16–22. [16] M. Qian, F.H. Froes, Titanium Powder Metallurgy: Science, Technology and Applications, Butterworth-Heinemann, Elsevier, 2015. [17] F. Arcella, F.H. (Sam) Froes, Production of titanium aerospace components from powder using laser forming, JOM 52 (5) (2000) 28. [18] A. Dehghan-Manshadi, M. Bermingham, M. Dargusch, D.H. St. John, M. Qian, Metal injection moulding of titanium and titanium alloys: challenges and recent development, Powder Technol. 319 (2017) 289–301. [19] F.H. (Sam) Froes, R.M. German, Titanium powder injection molding (PIM), Met. Powder Rep. 55 (6) (2000) 12. [20] R.M. German, Powder Injection Molding Design and Applications—Users Guide, Innovative Solutions, State College, PA, 2003. p. 5. [21] F.H. (Sam) Froes, Advances in titanium metal injection molding, in: M.N. Gungor, M. Ashraf Imam, F.H. (Sam) Froes (Eds.), Innovations in Titanium Technology, TMS (The Minerals, Metals & Materials Society), 2007, , pp. 157–166. [22] T. Kono, A. Horata, T. Kondo, Development of titanium & titanium alloy by metal injection molding process, Powder Powder Metall. 44 (11) (in Japanese).

Titanium background, alloying behavior and advanced fabrication techniques—An overview

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[23] W. Limberg, E. Aust, T. Ebel, R. Gerling, B. Oger, in: Metal injection molding of an advanced bone screw Ti-6Al-7Nb alloy powder, Euro, 2004. [24] Y. Itoh, T. Hankou, K. Sato, H. Miura, in: Improvement of ductility for injection moulding Ti-6Al-4V alloy, Euro, 2004. [25] H. Nakamura, T. Shimura, K. Nakabayashi, Process for production of Ti sintered compacts using the injection molding method, J. Jpn. Soc. Powder Powder Metall. 46 (8) (1999) (in Japanese). [26] A.T. Sidambe, Biocompatibility of advanced manufactured titanium implants—a review. Materials 7 (2014) 8168–8188, https://doi.org/10.3390/ma7128168. [27] K. Simmons, K. Scott Weil, E. Nyberg, Powder injection molding of titanium compounds, Ind. Heat. (2005) 43. [28] S. Guo, et al., Influence of sintering time on mechanical properties of Ti-6Al-4V compacts by metal injection molding, Rare Metal Mater. Eng. 34 (7) (2005) 33. [29] H. Wang, et al., Development of high density (99% +) powder injection molded titanium alloys, PIM Sci. Tech. Briefs I (5) (1999) 16. [30] K. Maekawa, et al., Effect of MIM process conditions on microstructures and mechanical properties of Ti-6Al-4V compacts, J. Japan Soc. PIM 46 (10) (1999) 1053. [31] K. Kusaka, et al., Tensile behavior of sintered Ti and Ti-6Al-4V alloy by MIM process, in: Advances in PIM and Particulate Materials, MPIF, Princeton, NJ, 1996, , pp. 29–127. [32] G. Wegmann, et al., Metal injection molding of titanium alloys for medical applications, Mater. Week (2000) 1. [33] B. Dutta, F.H. Froes, Additive manufacturing of titanium alloys, Adv. Mater. Process. (2014) 18–23. [34] B. Dutta, F.H. Froes, The additive manufacturing of titanium alloys, in: M. Qian, F.H. Froes (Eds.), Titanium Powder Metallurgy: Science, Technology and Applications, Butterworth-Heinemann, Elsevier, 2015 (Chapter 24). [35] M. Qian, W. Xu, M. Brant, H.P. Tang, Additive manufacturing and post-processing of Ti-6Al-4V for superior mechanical properties, MRS Bull. 41 (2016) 775–783. [36] H.P. Tang, Q.B. Wang, G.Y. Yang, J. Gu, N. Liu, L. Jia, M. Qian, A honeycomb-structured Ti-6Al-4V oil-gas separation rotor additively manufactured by selective electron beam melting for aero-engine applications, JOM 68 (3) (2016) 799–805. [37] B. Dutta, F.H. Froes, Additive Manufacturing of Titanium Alloys, Elsevier Publishing, 2016.

The molecular orbital approach and its application to biomedical titanium alloy design

1.3

M. Morinaga Toyota Physical and Chemical Research Institute, Nagoya, Japan

1.3.1

Introduction

Nowadays, a term for alloy design seems trite because it has spread wide in the field of metal science for a long period of time. But it was quite new in the beginning of the 1980s when the author began alloy design research. From a historical point of view, we started using the term alloy design once the PHACOMP method was developed in 1964. Following this method, the formation of brittle phases such as the σ phase and the μ phase was predicted in Ni-based superalloys [1,2]. However, in those days alloy design was still practically so difficult that most alloys were developed relying on previous experience and trial-and-error experiments. However, today materials informatics [3] are commonly used in every field of materials to save a lot of time and cost required for material development. As the author recalls from those days, it seems as if we were in another world. In this paper, a historical account of the development of the molecular orbital approach will be given, including not only Ti alloys but also Ni-based superalloys. This will show the whole feature of the theory of alloy design constructed on the basis of the molecular orbital method. Then, the phase stability and corrosion resistance of titanium alloys will be explained along this approach, because these properties are important for the design of titanium alloys for biomedical applications. Also, a concrete way of alloy design will be shown, using an example of high strength β-type Ti alloys. Then, the biomedical titanium alloys developed by this approach will be reviewed while touching on recent progress in this approach to biomedical alloy design.

1.3.2

Theory of alloy design

1.3.2.1 Alloying parameters Physical and chemical properties of metals and alloys are closely connected with their electronic state. Most commercially available alloys consist of multiple components. Some alloys consist of about 10 elements. For such complex alloys, the accurate calculation of the electronic state is too difficult to be performed in a reasonable manner. Titanium in Medical and Dental Applications. https://doi.org/10.1016/B978-0-12-812456-7.00003-2 © 2018 Elsevier Inc. All rights reserved.

40

Titanium in Medical and Dental Applications

Even if the calculation could be completed, it might be meaningless for practical alloy design because there are an infinite number of combinations of alloying species and compositions in multiple-component alloys. For practical alloy design, it is convenient to use alloying parameters that represent well the feature of each element in a mother metal. However, no alloying parameters have been found despite a long history of metallurgy and metal science. So, most alloy properties have been estimated by using the parameters assigned to pure metals (e.g., atomic radius, electronegativity, the electrons-per-atom-ratio (e/a)) [4,5]. In other words, the same value of the parameter is always used for the same element even if the alloy system is varied. As any alloying effects are not involved in the pure parameter, the prediction of alloy properties is inevitably poor along this approach. To solve this problem, we have to first determine alloying parameters theoretically using molecular orbital methods. The DV-Xα cluster method is used for this purpose. This is the first-principle molecular orbital calculation method. As compared with the conventional band calculation method, it is more suitable for the simulation of local electronic structures around alloying elements in metal. Further detailed explanation of the DV-Xα cluster method is given elsewhere [6,7]. Transition-metal based alloys (e.g., Fe alloys and Ti alloys) are very common in our daily life. For alloy design, it is very important to choose alloying parameters that match the use for alloys. For structural and biomedical use, two alloying parameters are essential to the design of transition metal-based alloys [8–10]. One alloying parameter is the bond order (hereafter referred to Bo). As shown in Fig. 1.3.1A, Bo shows the overlap population between the atomic orbital ϕΜ for alloying element (M) and ϕX for mother metal (X). As the overlap population increases, the covalent bond strength increases between the M and X atoms. When both M and X are transition metals, the d-d covalent bond is most dominant between them. The larger Bo value means the stronger chemical bond operating between M and X. Another alloying parameter is the d orbital energy level (hereafter referred to Md). As shown in Fig. 1.3.1B, the d orbitals of isolated M and X atoms are hybridized to form the bonding level at the lower energy and the antibonding level at the higher energy. When the d level of an isolated M atom is higher than that of an isolated X atom, charge transfer takes place from M to X so as to reduce energy. As a result, the effective charge becomes positive for M and negative for X. In this way the energy

Fig. 1.3.1 Alloying parameters, (A) bond order, Bo and (B) d-orbital level, Md. In (A), ϕM and ϕX are the atomic orbitals of M and X, and Bo is proportional to the overlap population between them.

f

fx

M

M Overlap integral =

(A)

X

f

M

M–d level

Charge transfer X–d level

f x dv (B)

The molecular orbital approach and its application to biomedical titanium alloy design

41

level controls the direction for charge transfer; hence it is a measure of the electronegativity. The lower electronegativity element has the higher d orbital energy level (Md). Also, it is noted that Md varies depending on the atomic radius. Any larger atom has the larger average radius of the d orbital. In this case, as the average distance increases between d electrons and the nucleus at the center, the attractive Coulomb force operating between them becomes weak. As a result, the d orbital level (Md) becomes high with increasing atomic radius. It is important to note here that Md correlates with both the atomic radius and the electronegativity, being classical parameters that have been used for treating the solid solubility problem of alloys [4,5]. Both Md and Bo are new alloying parameters first determined from the electronic structure calculation. As these are the electronic parameters, their magnitudes change following the order of elements in the periodic table. Recalling that most physical and chemical properties change following the periodic table, we know that these are very appropriate parameters for treating alloy properties. In addition, it is stressed that these two parameters represent the characteristic of the two-dimensional periodic table. The use of only one parameter (e.g., the Mendeleev number) never expresses any alloy property changes following the periodic table.

1.3.2.2 Molecular orbital calculation for nickel alloys This approach was first applied to the Ni-based superalloys, in which the Ni3Al (γ0 ) phase precipitates in the Ni (γ) matrix. The Ni3Al (γ0 ) phase is a strengthening phase and its volume fraction is as high as 60% in advanced Ni-based superalloys, which are commonly used for jet engines and gas turbines. As shown in Fig. 1.3.2A, Ni3Al has an L12-type structure [8]. By placing an Al atom at the center, the cluster model is made as shown in Fig. 1.3.2B. It consists of a central Al atom, 12 Ni atoms in the nearest neighbor, and 6 Al atoms in the second nearest neighbor. An alloying element, M, is substituted for a central Al atom. So, the cluster model is expressed as (MNi12Al6), and the 3d, 4d, and 5d transition metals are chosen for M. As M is surrounded by 12 Ni atoms, M is placed in a circumstance as in fcc Ni. The interatomic distances are set from the measured lattice parameter. The density of the states of electrons for pure Ni3Al is obtained from the cluster

Fig. 1.3.2 (A) Crystal structure of Ni3Al and (B) cluster model used in the calculation.

Ni AI

(A)

M

(B)

42

Titanium in Medical and Dental Applications

Fig. 1.3.3 Energy level structures of pure and alloyed Ni3Al with 3d transition elements. 16eg and 14t2g are d-orbital levels of alloying element, M.

calculation. The main features resemble the results of the band calculation in spite of a relatively small cluster model used in the present calculation [8]. The calculated level structures are shown in Fig. 1.3.3, where the Fermi energy level, Ef, of Ni3Al is set at zero and used as a reference. For a pure Ni3Al cluster, the levels of 13a1g to 15eg are derived mainly from the Ni 3d orbitals and form the Ni 3d band where the Ef lies, as indicated by an arrow. For the cluster alloyed with 3d transition metals (M), new energy levels originated mainly from the M-d orbitals appear above the Ef. For example, they are the 16eg and 14t2g levels drawn by dotted lines in the figure. Their height decreases monotonously with the atomic number of M. The average value of these two levels is defined as the Md level for M. Also, the bond order, Bo, between the Ni-d and M-d electrons is obtained for various M from the Mulliken population analysis [11]. The calculated values of Md and Bo are listed in Table 1.3.1 [12]. The bond order shows a maximum at the 6A group elements of Cr, Mo, and W. It is important to note here that the high Bo elements (e.g., Cr, Mo, Ta, W, and Re) are principal alloying elements in Ni-based superalloys [10]. The Bo is one of the indicators to choose alloying elements. For an alloy, the average values of Md and Bo are defined simply by taking the compositional average, and expressed as Md and Bo,

The molecular orbital approach and its application to biomedical titanium alloy design

43

Table 1.3.1 List of Md and Bo values for various alloying elements in Ni

3d

4d

5d

Md ¼

n X

Elements

Md (eV)

Bo

Ti V Cr Mn Fe Co Ni Cu Zr Nb Mo Tc Ru Rh Pd Ag Hf Ta W Re Os Ir Pt Au Al Si

2.271 1.543 1.142 0.957 0.858 0.777 0.717 0.615 2.944 2.117 1.550 1.191 1.006 0.898 0.779 0.659 3.020 2.224 1.655 1.267 1.063 0.907 0.764 0.627 1.900 1.900

1.098 1.141 1.278 1.001 0.857 0.697 0.514 0.272 1.479 1.594 1.611 1.535 1.314 1.068 0.751 0.391 1.518 1.670 1.730 1.692 1.500 1.256 0.920 0.528 0.533 0.589

Xi  ðMd Þi

(1.3.1)

Xi  ðBoÞi

(1.3.2)

i¼1

Bo ¼

n X i¼1

Here, Xi is the atomic fraction of component i in the alloy and (Md)i and (Bo)i are the Md and Bo values of component i. It is apparent that the Md represents a center of gravity in the d-band of the alloy, because Md is the compositional average of the d level of each alloying element in the alloy. If the alloy composition is known, both Md and Bo values are calculated readily using Eqs. (1.3.1), (1.3.2).

44

Titanium in Medical and Dental Applications

1.3.2.3 New PHPCOMP A solid solubility problem is very important in physical metallurgy. In the classical approach by Hume-Rothery and Darken-Gurry [4,5], this problem has been treated by using the atomic radius and the electronegativity. However, the problem still remains unknown when both solute and solvent atoms are transition metals. In 1964, the PHACOMP (PHAse COMPutation) method [1,2] was developed in the United States. This aimed to predict the solid solubility limit of elements in the Ni (γ) matrix and to avoid the formation of the brittle phases (e.g., the σ phase) in the γ matrix. The electron vacancy number, Nv, is used for this method. Here, Nv is the number of electron vacancies or holes existing above the Fermi energy level in the d band, and it is approximately expressed as Nv ¼ 10.66-e/a, where e/a is the electrons-per-atom ratio. For instance, the Nv value is 6.66 for the 4A group elements (i.e., e/a ¼ 4) of Cr, Mo, and W. However, Cr has a smaller atomic radius than Mo or W so that the atomic size concept is missing in the Nv parameter. Therefore, this method is contrary to the classical approach by Hume-Rothery and Darken-Gurry [4,5], and inevitably the prediction along this Nv method is poor. Nevertheless, PHACOMP has been used worldwide for the design and quality control of nickelbased superalloys. The author proposed a new PHACOMP in 1984 [9]. In this method, instead of the Nv parameter, the Md parameter is used for predicting the solid solubility limit. As mentioned before, the Md parameter is obtained for the first time from the calculation of the alloy cluster, which has the same chemical circumstance as in the fcc Ni alloy. Also the Md parameter is related closely to the electronegativity and the atomic radius of elements. In this sense, the Md parameter has the potential for predicting the solid solubility limit of alloys, both solute and solvent being transition metals. Two typical phase diagrams are shown in Fig. 1.3.4A for Ni-Co-Cr and Fig. 1.3.4B for Ni-Cr-Mo at 1477 K, in which the γ/γ + σ phase boundary is traced using the Co

Ni

1477 K R = 1.264 A

20

Md = 0.925

80

s Md = 0.925

a +g

(A)

mol% Cr

Mo

P

80

20

g + a 20

s

a

Ni 60

40

d a +s

40

g +s

60

40

g +s

80

20

mo l%

l%

g 60

60

g +P

Ni

Co

mo

60

80

l% mo

Nn = 2.49

l%

40

20

Nn = 2.49 g 40 + d

mo

Ni

R = 1.264 A

1477 K

g

80

Cr

Mo

(B)

20

40

60

80

Cr

mol% Cr

Fig. 1.3.4 Prediction of the γ/γ + σ phase boundary in ternary phase diagrams of (A) Ni-Co-Cr and (B) Ni-Cr-Mo.

The molecular orbital approach and its application to biomedical titanium alloy design

45

iso-Md line of 0.925. The unit of Md (eV) is omitted in the figure for simplicity. For comparison, both the iso-Nv line and iso-R line are drawn in each phase diagram. Here, Nv and R are the compositional averages of Nv and the atomic radius, R, of elements in the alloy, respectively. The iso-Nv line of Nv ¼ 2:49, which is often used for the prediction of the γ/γ + σ phase boundary, is far away from the boundary. On the other hand, the iso-R line is close to the boundary shown in (A), but it is far away from the boundary shown in (B). Compared to these, the iso-Md line is close to the boundary in both (A) and (B). The validity of this Md method has been confirmed through a series of examinations of more than 30 ternary phase diagrams [13–16]. We may even find some ambiguities involved in the experimental phase diagrams with the aid of the Md method [13–16]. In addition to ternary phase diagrams, commercially available alloys with multiple components are treatable by this Md method [9]. New PHACOMP has been applied successfully to the design of Ni-based superalloys [17–20].

1.3.3

Molecular orbital calculation and alloying parameters of titanium alloys

1.3.3.1 Molecular orbital calculation An allotropic transformation takes place between hcp Ti (α-Ti) and bcc Ti (β-Ti) around 1155 K in pure Ti. Both bcc and hcp cluster models are employed for the electronic structure calculation [21]. For example, the bcc cluster model is shown in Fig. 1.3.5A. It consists of a central Ti atom, eight Ti atoms in the nearest neighbor, and six Ti atoms in the second nearest neighbor. An alloying element, M, is substituted for a central Ti atom. So, the bcc cluster model is expressed as MTi14. The lattice parameter used is 0.3320 nm. On the other hand, the hcp cluster model is shown in Fig. 1.3.5B. It consists of a central Ti atom, 12 Ti atoms in the nearest neighbor, and six Ti atoms in the second nearest neighbor. An alloying element, M, is substituted for a central Ti atom. So, the

Fig. 1.3.5 Cluster models used for calculations, (A) bcc cluster and (B) hcp cluster.

Ti M

(A)

(B)

46

Titanium in Medical and Dental Applications

hcp cluster model is expressed as MTi18. The lattice parameters used are a ¼ 0.2950 nm and c ¼ 0.4683 nm. By using these cluster models, a series of calculations is carried out for various alloying elements, M, to determine alloying parameters relevant to titanium.

1.3.3.2 Alloying parameters In Fig. 1.3.6 the energy level structure is given of 3d alloying elements in bcc Ti [21,22]. For example, the result of pure Ti calculated by using the Ti15 cluster model (M ¼ Ti) is shown in the left side of the figure, where the Fermi energy level, Ef, is indicated by an arrow ( ). The levels of 18t1u to 5eu are derived mainly from Ti 3d orbitals, and form the Ti 3d band where the Ef lies. The lower energy levels of 17t1u to 15a1g and the higher energy levels of 13eg to 8t1g are composed of Ti 3d, 4 s, and 4p. Two energy levels, 16t2g and 13eg, existing around 2.5 eV are the d-levels of a central Ti atom in the cluster model shown in Fig. 1.3.5A. The fractions of the d component are about 40% in 16t2g and 59% in 13eg. In alloyed bcc Ti with 3d transition elements, the energy levels originated mainly from the M-d orbitals appear above the Ef. These M-d orbital levels are the eg and t2g levels indicated by dotted lines in Fig. 1.3.6, and their appearance is very similar to the results for the alloyed Ni3Al shown in Fig. 1.3.3. The height of these levels decreases

Fig. 1.3.6 Energy level structure of bcc Ti containing 3d transition elements.

The molecular orbital approach and its application to biomedical titanium alloy design

47

monotonously with the atomic number of M. The average value of these two levels is defined as Md for the Ti alloys. Another alloying parameter is the bond order (Bo). The bond order relevant to the d-d covalent bond is calculated following the Mulliken population analysis [11]. The results are shown in Fig. 1.3.7A for bcc Ti and in Fig. 1.3.8A for hcp Ti. In each figure, M-Ti and Ti-Ti are the bond order between M-d and Ti-3d electrons and between Ti-3d and Ti-3d electrons in a cluster, respectively. The total is the sum of them. The M-Ti bond order varies depending largely on the alloying element, M, in both bcc Ti and hcp Ti. There is a small peak near V, and then it decreases with the order of elements, M. Total bond order varies in a similar manner, as does the M-Ti bond order. The magnitude of the total bond order is smaller in 3d elements than in 4d and 5d elements. The trend of the bond order change with M resembles that between bcc Ti and hcp Ti. As described in the Ni-based superalloys [10], the high bond order elements are principal alloying elements in most of the alloys for structural applications. This is also the case in the titanium alloys because the principal alloying elements such as V, Cr, Zr, Nb, and Mo have the high bond order. In order to allow for the density and the specific strength of alloys, the bond order of each element is divided by its atomic weight. The results are given in Fig. 1.3.7B for bcc Ti and in Fig. 1.3.8B for hcp Ti. It is apparent that Fig. 1.3.7B is very similar to Fig. 1.3.8B. The ratio of bond order to atomic weight is large for Al, (Ti), and V. It may be said from this

3d

Bond order

3.0

Bond order/Atomic weight

Total Ti-Ti

M-Ti

0 0.10 0.08 0.06

Total

0.04 0.02 0

(B)

5d

2.0

1.0

(A)

4d

Ti-Ti M-Ti Al Si Sn Ti V CrMnFeCoNi Cu ZrNbMoHf Ta W M

Fig. 1.3.7 (A) Bond order and (B) ratio of bond order to atomic weight for M in bcc Ti.

48

Titanium in Medical and Dental Applications 5.0

Fig. 1.3.8 (A) Bond order and (B) ratio of bond order to atomic weight for M in hcp Ti.

3d

Bond order

4.0

5d

Total

3.0

Ti-Ti 2.0 1.0

(A)

4d

M-Ti

0.0

Bond order/Atomic weight

0.12 0.10 0.08

Total

0.06 0.04

Ti-Ti

0.02

M-Ti

0.00

(B)

Al Si Sn Ti V Cr MnFe CoNi CuZr NbMoHf Ta W M

result that the selection of alloying elements in the Ti-6Al-4V system is very reasonable. In Table 1.3.2, the Bo and Md parameters are listed for various alloying elements, M, in bcc Ti. Hereafter, the parameters obtained for bcc Ti are employed for the following analysis of alloy properties because these parameters appear to be relatively insensitive to the crystal structure [21,22].

1.3.4

Correlation of alloying parameters with alloy properties

1.3.4.1 Classification of binary phase diagrams Three typical phase diagrams of binary Ti-M alloys are illustrated in Fig. 1.3.9A–C, where M is a transition metal. Here, (A) is the α and β perfect solid solution type, (Β) is the β-isomorphous type, and (C) is the β-utectoid type. As shown in Fig. 1.3.9, these three types of phase diagrams can be distinguished clearly on the Bo versus Md diagram. The Ti-W phase diagram was previously grouped into (C) but according to a recent study it was revised and grouped into (B) [23], in agreement with the present classification.

Table 1.3.2

List of Md and Bo values for various alloying elements in bcc Ti

3d

Bo

Md (eV)

4d

Bo

Md (eV)

5d

Bo

Md (eV)

other

Bo

Md (eV)

Ti V Cr Mn Fe Co Ni Cu

2.790 2.805 2.779 2.723 2.651 2.529 2.412 2.114

2.447 1.872 1.478 1.194 0.969 0.807 0.724 0.567

Zr Nb Mo Tc Ru Rh Pd Ag

3.086 3.099 3.063 3.026 2.704 2.736 2.208 2.094

2.934 2.424 1.961 1.294 0.859 0.561 0.347 0.196

Hf Ta W Re Os Ir Pt Au

3.110 3.144 3.125 3.061 2.980 3.168 2.252 1.953

2.975 2.531 2.072 1.490 1.018 0.677 0.146 0.258

Al Si Sn

2.426 2.561 2.283

2.200 2.200 2.100

Fig. 1.3.9 Representation of binary Ti-M alloys on the Bo versus Md diagram.

Titanium in Medical and Dental Applications

Temp.

50

β α

α Ti

M% (A)

β

β α

α+β

M Ti

M% (B)

M Ti

M% (C)

M

3.5 (A) Type (B)

W

Ta

Hf

(C) 3.0

Mo

Nb

Zr

V Bo

Fe 2.5

Cr

TI

Mn

Co NI

Cu 2.0 0.0

1.0

2.0 Md (eV)

3.0

1.3.4.2 Classification of practically used alloys into α, α + β, and β-types We usually classify titanium alloys into the α, α + β, and β types from the phases existing in the alloy. About 40 practically used alloys are presented on the Bo  Md diagram shown in Fig. 1.3.10, where the Bo and Md values for each alloy are calculated from the alloy composition using Eqs. (1.3.1), (1.3.2) [21,22]. Here, the alloy compositions are denoted in mass%. In binary Ti-M alloys, the high-temperature β phase is retained metastably even at room temperature when the alloy composition exceeds a certain critical value. The critical value changes with M, and such data, (a)–(g), for various M are plotted on this diagram. It is evident from Fig. 1.3.10 that three-types of alloys are located separately on this diagram. For instance, the Ti-6Al-4V alloy (No. 9) exists in the α + β type field. Also, the Ti-8Mn alloy (No. 6) exists in the β field despite the α + β type alloy. However, this is not a contradiction because the Ti-8Mn is really a β type alloy, because the 8% Mn content in the alloy exceeds a critical value of 6% Mn, as shown in Fig. 1.3.10. But the alloy is heat treated in the (α + β) temperature range to improve the mechanical property, and incidentally grouped into the α + β type alloy. Even if the alloy type is

The molecular orbital approach and its application to biomedical titanium alloy design

1. Ti-4.5 Sn-11.5 Mo-6 Zr (Beta III) 2. Ti-3 AI-8 V-4 Zr-4 Mo-6 Cr (Beta C) 3. Ti-3 AI-8 V-8 Mo-2 Fe (8-8-2-3) b alloy 4. Ti-3 AI-13 V-11 Cr (13-11-3) a +b 5. Ti-15 Mo-5 Zr-3 AI 1 a 6. Ti-8 Mn (8 Mn) 5 a 7. Ti-6 AI-6 Mo-2 Sn-4 Zr (6-2-4-6) ec Pure Ti X 8. Ti-5 AI-2 Sn-2 Zr-4 Mo-4 Cr (Ti-17) 23 6d b 9. Ti-6 AI-4 V (6-4) f 10. Ti-6 AI-6 V-2 Sn (6-6-2) 8 11. Ti-2.25 AI-11 Sn-5 Zr-1 Mo-0.2 Si (IMI-679) 7 12 g 9 12. Ti-6 AI-0.5 Mo-5 Zr-0.2 Si (IMI-685) 1311 10 13. Ti-6 AI-2 Mo-2 Sn-4 Zr (6-2-4-2) 15 17 14. Ti-5 AI-6 Sn-2 Zr-1 Mo-0.2 Si (5621 S) 16 14 15. Ti-5 AI-2.5 Sn (A-110) 16. Ti-8 AI-1 V-1 Mo (8-1-1) 17. Ti-5.5 AI-3.5 Sn-3 Zr-0.3 Mo-1 Nb-0.3 Si 2.30 2.35 2.40 2.45 (IMI-829) Md (eV) Ti-M(mass%) (a) M=Mo(10%), (b) Fe(4%), (c) Cr(8%), (d) Mn(6%), (e) V(15%), (f) Co(6%), (g) Ni(8%)

2.84 2.82 2.80 Bo

51

2.78 4 2.76 2.74 2.72 2.25

Fig. 1.3.10 Grouping of commercial titanium alloys into the three types of α, α + β, and β alloys on the Bo  Md diagram.

uncertain, it is predictable by calculating the Bo and Md values from the composition and plotting the alloy location on the Bo  Md diagram shown in Fig. 1.3.10. No experiments are needed for this prediction. Furthermore, the alloying vector for binary Ti-M alloys is presented in Fig. 1.3.11 to understand the alloying behavior of respective elements on the Bo  Md diagram. The alloying vector starts at the position of pure Ti and ends at the position of Ti-10mass% M alloy. By comparing Fig. 1.3.10 with Fig. 1.3.11, we may see that, for example, the alloying vector for the Ti-Al alloy gets into the α-phase field as Fig. 1.3.11 Phase stability change with alloying elements. The vector represents the location of Ti-10mass% M alloy.

2.84 W Nb

Ta

Mo

2.82

Bo

β

2.80

V

2.78

Cr Mn Fe Co

Zr Hf Ti

2.76 Ni

α

2.74 Si

Cu

2.72 2.25

2.30

2.35

Sn

AI

2.40

Md (eV)

2.45

2.50

52

Titanium in Medical and Dental Applications

the Al content increases, indicating that Al is a α-stabilizing element. On the contrary for example, V, Nb, and Ta are the β-stabilizing elements because their vectors are directed toward the β-phase field. These results are in agreement with the well-known alloying behavior of elements in titanium.

1.3.4.3 Boundary between slip and twin deformation In the course of the deformation of β-type alloys, either the slip or the twin mechanism is operating, depending on the stability of the β-phase [24]. The slip mechanism is dominant when the alloy has the high β-phase stability whereas the twin mechanism is dominant when the alloy has the relatively low β-phase stability. The boundary compositions between the slip and twin mechanism are determined by examining deformation bands that appear around a Vickers indentation. As shown in Fig. 1.3.12, wavy slip bands appear when the slip mechanism is dominant whereas straight twin bands appear when the twin mechanism is dominant. This observation is carried out for a variety of binary alloys such as Ti-V, Ti-Cr, Ti-Mn, Ti-Fe, Ti-Co, Ti-Nb, Ti-Mo, Ti-W, and also for multiple-component Ti-Al-V-Cr-Mo-Zr alloys with various compositions. The results are summarized in the Bo  Md diagram shown in Fig. 1.3.12. It is noticed that most of the commercially available alloys exist along the slip/twin boundary. For instance, Ti-10%V-2%Fe-3%Al (10-2-3) is in the region of the twin mechanism, and Ti-15%V-3%Cr-3%Sn-3%Al (15-3-3-3) is in the region of the slip mechanism near the slip/twin boundary.

2.84

Twin Twin>Slip 15-5 (Twin)

Twin=Slip

2.82

Twin Al > Sn > Zr. Special alloying behavior emerges by the addition of these elements into the β-phase alloys [31].

1.3.5.4.3 Superelasticity and shape memory effect For Ti-Nb binary alloys, superelasticity appears in the least stable β-phase alloy (i.e., Ti-(40–42) Nb (mass%)) [43], in which the reversible transformation takes place between the β-phase and the martensite at RT. Because the energy difference between the β-phase and the martensite α00 phase is very small in such a least stable β-phase alloy. The shape memory effect emerges in the lower Nb-content alloy (i.e.,Ti(35-40) Nb) [43], in which the Ms temperature is close to RT. However, neither the superelasticity nor the shape memory effect appears in the more stable β-phase alloys (>42Nb), in which the Ms temperature is well below RT. For a lower Nb-content alloy (i.e., Ti-30Nb) that is located in the martensite zone, neither the superelasticity nor the shape memory effect emerges. However, as indicated by alloy number ⑬ in Fig. 1.3.16, the addition of 1% Fe to this Ti-30Nb alloy moves the alloy location to the least stable β-phase zone so it exhibits the superelasticity. A compositional range of the alloys where either the superelasticity or the shape memory effect emerges is modified by the addition of O, Al, Sn, and Zr. For instance, as mentioned above, the shape memory effect emerges in a Ti-38.5Nb alloy but the further addition of 1.5% Al to this alloy yields a 4.3% superelastic strain [44]. Such a variation in the alloy property is also seen in Ti-36Nb by changing the O content in it. Namely, the shape memory effect still emerges in Ti-36Nb-(0–0.12) O alloys but the further addition of O causes the superelasticity in the Ti-36Nb-(0.22–0.43) O alloys, in which the least stable single β-phase exists at RT [36]. Thus, the emergence of the superelasticity is further promoted in the O-containing Ti-Nb alloys than in the O-free Ti-Nb alloys [36]. This is also true in the multiple-component Ti-5Zr-pNb-30Ta-0.23O alloys (p ¼ 20, 25, 30, 35, and 40) [39]. Neither phenomenon is observed in the stable β-phase alloys (i.e., 35Nb). However, as shown in Fig. 1.3.17, the superelasticity emerges in the least stable β-phase alloys (i.e., 30Nb). The stress-induced martensitic transformation, which is associated with the shape memory effect, is observed in the lower Nb content alloys (i.e., 20Nb and 25Nb) [40]. It is noted here that the least stable β-phase alloy (i.e., 30Nb) has the lowest Young’s modulus, as explained before. Thus, the β-phase stability of the alloys correlates well with the alloy properties such as the low Young’s modulus, the superelasticity, and the shape memory effect, and such a correlation is traceable on the Bo  Md diagram. The special alloying of O, Al, Sn, and Zr induces a remarkable phase boundary shift, which will give us plenty of scope for the modification of the phase stability and hence the alloy properties in a reasonable way.

62

1.3.6

Titanium in Medical and Dental Applications

Conclusion

The alloy design theory is constructed not only for the fundamental understanding of alloys but also for practical alloy design without repeating trial-and-error experiments [45]. In particular, this theory is useful for the design of titanium alloys because the Bo  Md diagram represents well the phase constitution in the alloys. Such knowledge of the phase stability presented on the Bo  Md diagram is important in predicting various alloy properties such as the mechanical strength, corrosion resistance, Young’s modulus, superelasticity, and shape memory effect. Reminding that all these properties are essential to the biomedical titanium alloys, it is greatly expected that the present theory works well in practice in the course of alloy design for biomedical applications.

Acknowledgments This research was supported by the Grant-in-Aid for Scientific Research from the Ministry of Education, Culture, Sports, Science and Technology of Japan, and also from the Japan Society for the Promotion of Science (16K06711).

References [1] W.J. Boesch, J.S. Slaney, Preventing sigma phase embrittlement in nickel base superalloys, Met. Progr. 86 (1964) 109–111. [2] C.S. Barrett, Some industrial alloying practice and its basis, J. Inst. Met. 100 (1972) 65–73. [3] For example, http://www.jst.go.jp/crds/pdf/2015/FR/CRDS-FY2015-FR-05.pdf. [4] W. Hume-Rothery, G.V. Raynor, Structure of Metals and Alloys, fourth ed., Institute of Metals, London, 1962. [5] L. Darken, R.W. Gurry, Physical Chemisry of Metals, McGraw-Hill, New York, 1953. [6] H. Adachi, M. Tsukada, C. Satoko, Discrete variational Xα calculations. 1. Application to metal clusters, J. Phys. Soc. Jpn. 45 (1978) 875–883. [7] H. Adachi, Introduction to Quantum Materials Chemistry—Approach from DV-Xα Method, Sankyo Publishing Co. Ltd., Tokyo, 1991 (in Japanese). [8] M. Morinaga, N. Yukawa, H. Adachi, Alloying effect on the electronic structure of Ni3Al(γ0 ), J. Phys. Soc. Jpn. 53 (1984) 653–663. [9] M. Morinaga, N. Yukawa, H. Ezaki, H. Adachi, New PHACOMP and Its Applications to Alloy Design, Superalloys 1984, Ed. by M. Gell et al., The Metallurgical Society of AIME, Warrendale, Pennsylvania, 1984, pp. 523–532. [10] M. Morinaga, Alloy design based on molecular orbital method, Mater. Trans. 57 (3) (2016) 213–226. [11] R.S. Mulliken, Electronic population analysis on LCAO-MO molecular wave functions. 1.,2.,3.,4, J. Chem. Phys. 23 (1955) 1833–1840. pp. 1841–1846, 2338-2342, 2343–2346. [12] M. Morinaga, Y. Murata, H. Yukawa, Recent Progress in the New PHACOMP Approach, Materials Design Approaches and Experiences, ed. by J.-C. Zhao, M. Fahrmann and T. M. Pollock, TMS, Warrendale, Pennsylvania, 2001, pp. 15–27. [13] M. Morinaga, N. Yukawa, H. Ezaki, H. Adachi, Solid solubilities in transition-metal-based f.C.C. alloys, Philos. Mag. A 51 (1985) 223–246.

The molecular orbital approach and its application to biomedical titanium alloy design

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[14] M. Morinaga, H. Ezaki, H. Adachi, Solid solubilities in nickel-based f.C.C. alloys, Philos. Mag. A 51 (1985) 247–252. [15] H. Ezaki, M. Morinaga, N. Yukawa, H. Adachi, Prediction of the occurrence of the σ phase in Fe-Cr-Ni alloys, Philos. Mag. A 53 (1986) 709–716. [16] H. Ezaki, M. Morinaga, N. Yukawa, H. Adachi, Solid solubilities of molybdenum and tungsten in FCC transition-metal alloys, Philos. Mag. A 65 (1992) 1249–1260. [17] P. Caron, T. Khan, Design of Superalloys for Single Crystal Blade Applications: A 20-Year Experience, Materials Design Approaches and Experiences, ed. by J.-C. Zhao, M. Fahrmann and T. M. Pollock, TMS, Warrendale, Pennsylvania, 2001, pp. 1–14. [18] N. Yukawa, M. Morinaga, Y. Murata, H. Ezaki, S. Inoue, in: D.N. Duhl et al., (Ed.), High Performance Single Crystal Superalloys Developed by the d-Electrons Concept, Superalloys, 1988 The Metallurgical Society of AIME, Warrendale, Pennsylvania, 1988, pp. 225–234. [19] K. Matsugi, Y. Murata, M. Morinaga, N. Yukawa, Realistic Advancement for Nickel Based Single Crystal Superalloys by the d-Electrons Concept, Superalloys 1992, Ed. By S. D. Antolovich et al., TMS, Warrendale, Pennsylvania, 1992, pp. 307–316. [20] R. Hashizume, A. Yoshinari, T. Kiyono, Y. Murata, M. Morinaga, K.A. Green et al., (Ed.), Development of Ni-Based Single Crystal Superalloys for Power-Generation Gas Turbines, Superalloys 2004, TMS, Warrendale, Pennsylvania, 2004, pp. 53–62. [21] M. Morinaga, N. Yukawa, H. Adachi, Electronic structure and phase stability of titanium alloys, Tetsu Hagane 72 (1986) 555–562 (in Japanese). [22] M. Morinaga, N. Yukawa, T. Maya, K. Sone, H. Adachi, in: Theoretical design of titanium alloys, Proc. Sixth World Conf. on Titanium, Cannes, France, June 6–9, 1988, Societe Franc¸aise de Metallurgie, Paris, 1988, pp. 1601–1606. [23] H. Okamoto, Phase Diagrams for Binary Alloys, second ed., ASM International Metals Park, OH, 2010. [24] M. Morinaga, M. Kato, T. Kimura, M. Fukumoto, I. Harada, K. Kubo, in: Theoretical design of β-type titanium alloys, Proc. 7th World Conference on Titanium, San Diego, California, June 29–July 2, 1992, TMS, Warrendale, Pennsylvania, 1992, pp. 217–224. [25] M. Morishita, Y. Ashida, M. Chikuda, M. Morinaga, N. Yukawa, H. Adachi, Active corrosion rate for Ti-based alloys in aqueous corrosion and its correlation with the bond order obtained by electron theory, ISIJ Int. 31 (1991) 890–896. [26] M. Morishita, M. Chikuda, Y. Ashida, M. Morinaga, N. Yukawa, H. Adachi, Electronic states of the cathodes of titanium-based alloys in aqueous corrosion, Mater. Trans. JIM 32 (1991) 264–271. [27] Y. Okazaki, A. Ito, T. Tateishi, Y. Ito, Effect of alloying elements on anodic polarization properties of titanium alloys in acidic solutions, Mater. Trans. JIM 35 (1) (1994) 58–66. [28] M. Semlitsch, F. Stab, H. Webber, Titanium-aluminium-niobium alloy, development for biocompatible, high strength surgical implants, Biomed. Tech. 30 (1985) 334–339. [29] D. Kuroda, M. Niinomi, M. Morinaga, Y. Kato, T. Yashiro, Design and mechanical properties of new β type titanium alloys for implant materials, Mater. Sci. Eng. A 243 (1998) 244–249. [30] T. Saito, T. Furuta, J.H. Hwang, S. Kuramoto, K. Nishino, N. Suzuki, et al., Multifunctional alloys obtained via a dislocation-free plastic deformation mechanism, Science 300 (2003) 464–467. [31] M. Abdel-Hady, K. Hinoshita, M. Morinaga, General approach to phase stability and elastic properties of β-type Ti-alloys using electronic parameters, Scr. Mater. 55 (2006) 477–480. [32] H. Ikehata, N. Nagasako, T. Furura, A. Fukumoto, K. Miwa, T. Saito, First-principles calculations for development of low elastic modulus Ti alloys, Phys. Rev. B 70 (2004) 174113.

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[33] L.A. Matlakhova, A.N. Matlakhova, S.N. Monteiro, S.G. Fedotov, B.A. Goncharenko, Properties and structural characteristics of Ti-Nb-al alloys, Mater. Sci. Eng. A 393 (2005) 320–326. [34] Y.L. Zhou, M. Niinomi, T. Akahori, Effects of ta content on Young’s modulus and tensile properties of binary Ti-ta alloys for biomedical applications, Mater. Sci. Eng. A 371 (2004) 283–290. [35] T. Ozaki, H. Matsumoto, S. Watanabe, S. Hanada, Beta Ti alloys with low young’s modulus, Mater. Trans. 45 (8) (2004) 2776–2779. [36] J.I. Kim, H.Y. Kim, H. Hosoda, S. Miyazaki, Shape memory behavior of Ti-22Nb(0.5–2.0)O (at%) biomedical alloys, Mater. Trans. 46 (4) (2005) 852–857. [37] J.I. Qazi, B. Marquardt, H.J. Rack, High-strength metastable beta-titanium alloys for biomedical applications, JOM 56 (11) (2004) 49–51. [38] S. Ishiyama, S. Hanada, O. Izumi, Effect of Zr, Sn and al additions of deformation mode and Beta phase-stability of metastable beta Ti alloys, ISIJ Int. 31 (1991) 807–813. [39] N. Sakaguchi, M. Niinomi, T. Akahori, T. Saito, T.Furuta, Effects of alloying elements on elastic modulus of Ti-Nb-ta-Zr system alloy for biomedical applications, Mater. Sci. Forum 449–4 (2004) 1269–1272. [40] N. Sakaguchi, M. Niinomi, T. Akahori, Tensile deformation behavior of Ti-Nb-ta-Zr biomedical alloys, Mater. Trans. 45 (4) (2004) 1113–1119. [41] M. Tahara, H.Y. Kim, T. Imamura, H. Hosoda, S. Miyazaki, Lattice modulation and superelaticity in oxgen-added beta-Ti alloys, Acta Mater. 59 (2011) 6208–6218. [42] J.I. Kim, H.Y. Kim, T. Inamura, H. Hosoda, S. Miyazaki, Shape memory characteristics of Ti-22Nb-(2–8)Zr(at%) biomedical alloys, Mater. Sci. Eng. A 403 (2005) 334–339. [43] H.Y. Kim, H. Satoru, J.I. Kim, H. Hosoda, S. Mitazaki, Mechanical properties and shape memory behavior of Ti-Nb alloys, Mater. Trans. 45 (7) (2004) 2443–2448. [44] Y. Fukui, T. Inamura, H. Hosoda, K. Walashima, S. Miyazaki, Mechanical properties of a Ti-Nb-Al shape memory alloy, Mater. Trans. 45 (4) (2004) 1077–1082. [45] M. Morinaga, Y. Murata, H. Yukawa, Molecular orbital approach to alloy design, Applied Computational Materials Modeling-Theory, Simulation and Experiment, ed. by G. Bozzolo et al., Springer, New York, 2007, pp. 255–306.

1.4

Titanium and titanium alloys: Materials, review of processes for orthopedics and a focus on a proprietary approach to producing cannulated bars for screws and nails for trauma F. Ory, J.L. Fraysse Forecreu SAS, Malicorne, France

The use of titanium in medical applications started in the 1950s and has been widely used ever because due to its biocompatibility with the human body as well as many other advantages. The use of titanium in joints such as the hip and knee prostheses is well known, as is its application in trauma plates, screws, and nails. However, titanium and titanium alloys may allow for the future development of bioactive implants, external prostheses, and instrumentation. Advantages of titanium or titanium alloys for biomedical applications: l

l

l

l

l

l

l

Biocompatibility. Stronger than bone [80–120 MPa], stainless steels, and other implant materials. High fatigue strength [50% higher than cast Co-Cr alloy]. High resistance to impact loading. Good corrosion resistance. Lower elastic modulus than steel or cobalt alloys. Low density.

1.4.1

General processes for titanium alloys: From ore to bar material

The processing of titanium mill products occurs in the following major steps: reduction of titanium ore into «sponge», metallic porous form; melting of sponge, or sponge plus master alloy to form an ingot; primary fabrication, where an ingot is forged into general mill products such as billets or slabs; and possibly rolled into semi-products such as plate, sheet, strip, bar/rod/wire or tube (Figs. 1.4.1–1.4.5). Typically, the ingot is converted into a bloom, which in turn is rolled to billet on a reversing mill. After billet preparation, the billet is rolled to bar and rod (Figs. 1.4.6–1.4.8). Titanium in Medical and Dental Applications. https://doi.org/10.1016/B978-0-12-812456-7.00004-4 © 2018 Elsevier Inc. All rights reserved.

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Fig. 1.4.1 Hip stem systems (Timet). How is titanium made ? Ore to metal • Rutile (TiO2) to TiCl4 • TiCl4 + 2Mg = Ti + MgCl2

Thermomechanical processing • Forging • Rolling: Plate/Sheet/Bar

Melting

Fabrication

• VAR • Electron beam cold hearth • Plasma cold hearth

• Investment casting • Forging/ Machining • Extrusion

Fig. 1.4.2 Timet.

1.4.2

Families of titanium and titanium alloys for orthopedics

1.4.2.1 Further processing of titanium alloys to near net shape 1.4.2.1.1 Forging Reconstructive implants made of metals can be either machined from solid bars or obtained from castings, but many are made from near-net forging. The orthopedic industry has been a driving force in improving forging technologies. Forging is one

Titanium and titanium alloys

Fig. 1.4.3 Rutile and titanium sponge (Timet).

Fig. 1.4.4 Titanium ingot (Timet).

Fig. 1.4.5 Titanium ingot (Timet).

67

68

Titanium in Medical and Dental Applications

Ingot

Blooms

Unconditioned billet

Ground bar and custom tolerance bar

Hot rolled bar

Precision shapes

Lubed fastener wire

Fine wire

Conditioned billet

Hot rolled coil

Seam-free coil

Weld wire

Fig. 1.4.6 Processes from ingot to wire/bar/shapes (Dynamet).

Fig. 1.4.7 Ti alloy profiles (Dynamet).

Shaped wire

Ground bar, custom tolerance bar, precision bar and ultrabar∗

Titanium and titanium alloys

69

Fig. 1.4.8 Coiling Ti alloy (Dynamet).

of the oldest metal-forming processes in the world, dating back to 4500 BCE. During the 10th century, forging was a primary process for making weapons such as swords. The procedure continues to be used to create high performance, high-strength products today (Figs. 1.4.9–1.4.11). Net shape forging capabilities impart a tighter and more uniform fine grain structure while increasing tensile strength and yield strength. The result is a stronger part strength than those produced by any other metalworking process, due to forging’s continuous directional grain flow.

1.4.2.1.2 Investment casting The precision casting by the lost wax process, called investment casting, dates back to 1000 BCE in Mesopotamia for art statues. The idea is to create wax patterns, which look like the metal part to manufacture, assemble them, coat them with multiple dips in liquid ceramics, and then spray them with refractory sand. By melting the wax, we obtain an empty cluster that will be poured with molten metal (Figs. 1.4.12 and 1.4.13). Further processing includes ceramic removal, dismantling of assembly, finishing, heat treatments, and nondestructive controls such as x-ray inspection or fluorescent die penetrant visual tests.

1.4.2.1.3 Additive manufacturing Additive manufacturing (AM) technology fundamentally differs from traditional material removal processes like milling, turning, and spark erosion, because this technology family creates metal components by incremental addition of material instead of removal of chips. Starting from a three-dimensional CAD representation of an

70

Designation ASTM

F67

ISO

5832-2 CP Grade2 T40

5832-3

F1472

5832-3

F1295

5832-11

Ti-6AI4V Grade23

N

≤ 0.03

C

H

≤ 0.08 ≤ 0.0125

Fe

≤ 0.2

O

≤ 0.25

Ti

Others

UTS -KSI (N/MM2)

YS -KSI (N/MM2)

Application

Comment

44(300)

Dental implants CMF and screws

Low mechanical constraints and appreciated flexibility deformability High purity for maximum biocompatibility Higher strength alloys than grade 2

balance

/

131(900)

116(800)

Reconstruction : THR, shoulder Trauma : screws, nails, plates, spinal

51(355)

≤ 0.05

≤ 0.08

≤ 0.012

≤ 0.25

≤ 0.13

balance

AI6 V4

Ti-6AI4V non-ELI ≤ 0.05 Grade 5

≤ 0.08

≤ 0.009

≤ 0.3

≤ 0.2

balance

AI6 V4

135(930)

120(830)

Reconstruction : THR, shoulder Trauma : screws, nails, plates, spinal

10% increased strength

Ti-6AI7Nb

≤ 0.05

≤ 0.08

≤ 0.02

≤ 0.25

≤ 0.20

balance

Ta ≤ 0.5 AI6 Nb 7

145(1000)

131(900)

THR, Trauma

Same strength as Ti6AI4V ELI

Ti-13Nb-13Zr

≤ 0.05

≤ 0.08

≤ 0.012

≤ 0.25

≤ 0.15

balance

Nb13 Zr13

125(860)

105(725)

Trauma

/

Ti-12Mo-6Zr-2Fe ≤ 0.05

≤ 0.05

≤ 0.02

≤ 2.5

≤ 0.28

balance

Mo12 Zr6

135(932)

130(897)

Trauma

/

F2066

Ti-15Mo

≤ 0.05

≤ 0.1

≤ 0.015

≤ 0.10

≤ 0.20

balance

Mo15

100(690)

70(483)

Nail, Trauma

/

F2146

Ti-3AI-2.5V

≤ 0.02

≤ 0.05

≤ 0.015

≤0.30

≤ 0.12

balance

AI3 V2.5

Nail, Trauma

Low strength

F1813

Fig. 1.4.9 Families of titanium and titanium alloys.

A 90(621) CW A 75(517) CW 125(862) 105(724)

Titanium in Medical and Dental Applications

F136

F1713

Common

Chemical analysis

Fig. 1.4.10 Hip stem forging (SMB Medical, Switzerland).

Fig. 1.4.11 Die penetrant (SMB Medical, Switzerland).

Fig. 1.4.12 Investment casting. From MediMet GmbH Germany.

72

Titanium in Medical and Dental Applications

Fig. 1.4.13 Total knee joint. From MediMet GmbH Germany. Laser Mirror scanner XY deflection

f-0 lens Protective atmosphere

Roller/scraper

Y

X Feed container

Z

Base plate Build cylinder Overflow container

Fig. 1.4.14 Schematic illustration of the selective laser melting process (top).

object, the object is virtually “sliced” into a set of two- dimensional layers. These layers are then successively fused and consolidated on top of each other to recreate the three-dimensional object (Figs. 1.4.14–1.4.16). AM technologies, which have existed for a few decades, were originally developed as plastic prototyping technologies. Over the last decade, metal AM technologies were developed in order to benefit from AM advantages for manufacturing metallic components. Selective laser melting (SLM) and electron beam melting are the best-established AM technologies for high-quality metal manufacturing. In these processes, thin layers of metal powder (typically 20–40 μm) are spread by a powder deposition system. Next, a focused laser beam is used to scan a two-dimensional layer of the three-dimensional component. This process repeats until the whole threedimensional object has been “printed.”

Titanium and titanium alloys

73

Fig. 1.4.15 SLM production of a dental implant suprastructure in Ti-6Al-4V (right) (3D Systems, Belgium).

Fig. 1.4.16 Serial-produced hip implant with integrated bone scaffold, improving both initial and long-term stability (3D Systems, Belgium).

Due to the unlimited geometric freedom, patient-specific implants are perfectly suited to be produced with SLM technology. Besides custom implants, these technologies also introduce new possibilities for manufacturing standard implants and instrumentation. Orthopedic implants, for example, can be designed with integrated porosity regions for improved osseointegration. As an illustration. Fig. 1.4.3 shows a titanium femur incorporating four different scaffold designs and various degrees of surface finish.

74

Titanium in Medical and Dental Applications

1.4.2.1.4 Machining orthopedics Medical products and technologies are contributing more and more to human health, and are promoting mobility and vitality well into old age. Increasing demand and high standards imposed on the quality of medical products characterize the development of a sector that is growing in spite of the economic crisis. Sustained progress in medical engineering regularly places new demands on the machines that are used to manufacture high-accuracy products. Fifty years ago, the idea of implants and artificial joints in surgery was still revolutionary; today it is a fixed part of everyday medicine. Complex fractures are corrected with bone screws and plates, and, these days, worn hip joints are easy to replace. However, products of appropriate high quality are required for such surgical measures. These not only necessitate the use of superior materials that are usually difficult to machine but also place the highest demands on precision, surface quality, and dimensional accuracy of the medical products (Fig. 1.4.17). The example of artificial joints shows that efficient machining can only be carried out on high- performance milling machines. Superior materials such as titanium and also special materials such as zirconium, ceramics, and special plastics are reason enough for using modern manufacturing methods such as HSC milling. A similar argument applies to implants, which in traumatology are often only inserted temporarily and removed once the bone has healed. In this case, the highest demands are placed on the surface finish so that the implant does not adhere to the bone (Figs. 1.4.18 and 1.4.19). Fig. 1.4.17 Machining knee joint (DMG Mori, Germany).

Titanium and titanium alloys

75

Fig. 1.4.18 Spinal pedicle screws (Sandvik Coromant, Sweden).

Fig. 1.4.19 Hip cup (DMG Mori, Germany).

1.4.3

Proprietary approach to producing cannulated bars for screws and nails for trauma

1.4.3.1 Focus on cannulated 1.4.3.1.1 Minimally invasive Kirschner wire-guiding technique Dr. Kirschner first introduced the guidewire technique back in 1910 to do fracture fixation with an aiming device for trauma. Since then, this surgical technique has been extended with success in arthroscopy, sports medicine, small fragments, extremities, and spinal treatments.

76

Titanium in Medical and Dental Applications

Fig. 1.4.20 (Top) K-wire orthopedic surgery.

This guidewire is made of a stiff stainless steel wire usually 316L, aged 630, and with a shape memory alloy of Nitinol or CoCrMo with high yield strength. Usual diameters range from 0.6 to 3.2 mm depending on the size of the implants. It has a sharp or threaded point that can enter the bone with limited trauma. The idea is to position this wire in the bone and to use it as a guide for the next steps that are usually drilling, depth control, tapping, reaming, etc. Thus the orthopedic surgeon can, while maintaining the K-wire in the bone, insert the cannulated screw with a driver, cannulated as well. Using this technique guarantees precise and secured access to rare bone materials for fracture fixation (Figs. 1.4.20 and 1.4.21).

1.4.3.1.2 Cannulated instruments and implants Demand from surgeons calls for more minimally invasive surgery (MIS) systems. The requirement is then to offer instruments, nails, and screws with a central hole. Trauma products are usually machined on CNC lathes with bar feeders. Creating a concentric longitudinal hole is not an easy technique and several technologies can be used, including drilling/boring, using thick wall tubing, and proprietary precannulated bars (Figs. 1.4.22–1.4.24).

Titanium and titanium alloys

77

Fig. 1.4.21 (Bottom) X-ray compression hip screw (Forecreu, France).

Fig. 1.4.22 Cannulated bars, screws, and nails (Forecreu, France).

1.4.3.1.3 Tubing versus cannulated bars Conventional tubes with a thin or medium thick wall are produced according to the following steps. A billet (Fig. 1.4.25) is drilled through, heated to forging temperatures, and extruded. A needle inserted from the ram to the nose of the billet creates a tube hollow as the material flows between the needle and the die.

78

Titanium in Medical and Dental Applications

Fig. 1.4.23 X-ray small fragments (Forecreu, France).

DHS DCS screw

Spinal screw

Femoral, tibia nails

Small fragments

Cannulated screws

Fig. 1.4.24 Cannulated applications (Forecreu, France).

Titanium and titanium alloys

79

Billet

Extrusion

Tube

Billet in loading position Container Stem

Tube sinking

Die

Steel backing Carbide insert

Glass film Glass disc

Mandrel

Billet, 1200°C

3 – 5 m/s Lining

Extrusion in progress

Rod drawing

Floating plug drawing

Steel backing Carbide insert Tube

Steel backing Carbide insert

Carbide plug

And / Or

Mandrel

Fig. 1.4.25 Conventional tubing process (Forecreu, France).

Then several drawing techniques are combined and applied on the tube hollows as shown. The limitations are the surface geometry from the sinking process and the potential risk of capturing lubricant with the carbide plug or hard-mandrel process. As annealing steps are necessary to reduce the section, leftover lubricant traces can become intergranular corrosion. This process becomes difficult as the ID gets smaller and the wall thicker.

1.4.3.1.4 Manufacturing cannulated A special process has been designed to process material as a solid by plugging inserts into the starting material as shown in the following processing route. Properties of this process are: l

l

l

l

Conservation of homotetical ratio. No sinking during tube drawing, avoiding surface defects. No contamination risk by lubricants during annealing or heat treatments. No limitation in ID sizing to fit the smallest K-wires such as 0.02300 (0.6 mm) to 0.03100 (0.8 mm).

Using bimetal extrusion, rolling and drawing hot forming techniques. Starting with a 100/120 mm (4–500 ) round billet, approx. 400–600 mm long, the cylinder is deep-hole drilled. Then an insert, which has previously been coated with a stop-off layer, is introduced. The first step consists of the extrusion of the cylinder at the forging temperature. The extrusion pushes the hot billet through a die at an average speed of 2–4 m/s with up to 25–30 times elongation. Further processing includes rolling with a triangular deformation and finishing by warm drawing/calibrating.

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Titanium in Medical and Dental Applications

To obtain the final hole, the inserted material core that has become a wire must be removed. Wires are extracted by pulling, using the properties of ductility of the inserted (Figs. 1.4.26–1.4.29). l

Focus on core material

Composite hot extrusion requires that the insert flow smoothly through the die. Depending on forging temperature, the idea was to develop a combination of

Material control & testing To ASTM & ISO

Material inventory

Insertion of mandrel in the hole

Cutting of billets

Boring of the hole Insert / core

Hot extrusion

Ram

Semi-finished extrusions

Fmax ≈1310 T

Calibration Drawing effort Fmax ≈ 20 T

Kocks rolling mill Die

Bar

Wire removal

Centerless grinding

Dimensional testing

Hole cleaning

Final mechanical and metallographic testing to ASTM & ISO

Fig. 1.4.26 Processing flow chart (Forecreu, France).

Straightening

Titanium and titanium alloys

Fig. 1.4.27 (Top) 1310 tons horizontal press LOEWY +3 cells  350 kW BANYARD induction heating (Forecreu, France).

Fig. 1.4.28 Extrusion operation (Forecreu, France).

Fig. 1.4.29 (Bottom) 1310 tons horizontal press LOEWY +3 cells  350 kW BANYARD induction heating (Forecreu, France).

81

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Titanium in Medical and Dental Applications

Fig. 1.4.30 Insertion of core (Forecreu, France).

core-billet material that will have the same strength at extrusion temperatures, a core that would reduce proportionally. High ductility was required for final core/wire extraction/pullout. Mandrel materials were then developed to fit this need. The final decision called for a highly ductile self-reducing lost core. The metallurgy of it is proprietary (Fig. 1.4.30). l

Focus on choice of coating films

The prerequisite of the medical orthopedics application and hot-forging processes calls for an inert material, stop-off membrane, and anti-welding prevention, however removable and biocompatible, among multiple features. The challenge was to make sure the coating would create the barrier during lengthy heat treatments against any metallurgical transfer-generating welding points. l

Focus on extrusion

It is widely used in ferrous and nonferrous metal for processing into tube hollows, profiles, or composite and solid material (Fig. 1.4.31). Lubricant plays an active role in the process itself. Up to 600°C, the usual lubricants are graphite-based coating while steel or high temperature extrusions above 900°C are glass-based. Both the billet and/or container can be lubricated. Titanium alloy billets are coated prior to heating, then rolled on a bed of glass. The container is also coated. Any rupture in the coating during extrusion will jeopardize the symmetry, as shown in this figure. Extrusion proves to be the only forging process that is capable of maintaining a proper positioning of the plugged hole.

Fig. 1.4.31 Extrusion (Forecreu, France).

Titanium and titanium alloys

83

The control of the melting temperature of glass is therefore critical, as 900°C extrusion or 950°C will require different types of glass (Figs. 1.4.32 and 1.4.33). l

Flow of extrusion

When introduced in the container, the coated billet will be pushed with a force of typically 1300 t (200 athm), compressed initially to fill the container. The flow will then be constant. Figs. 1.4.34 and 1.4.35 show the initial peak of pressure and then a somewhat constant pressure and speed up to 4 m/s, until it stops. The control of the combination of parameters (temperature, speed, and pressure) is the key to proper concentricity. l

Focus on drawing

Once the extruded composite bar is cleared for the next step (figure image recognition analysis), the bar end is made thinner with a swage and then drawn through a die, thus reducing the diameter. The composite bar is reduced as a solid. Depending on the alloy, the temperature will be adjusted from cold to mild and warm. Lubricant will

Fig. 1.4.32 Deformation after extrusion: (A) with lubrication, (B) cleaned (Forecreu, France).

Fig. 1.4.33 Flow of composite material during extrusion (Forecreu, France).

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Titanium in Medical and Dental Applications

TA6V/20.1/3.56 950° 11 450 mm

800

Force x Vitesse (kN. m/s)

F.V. 40

F.V. 41

F.V.

F.V. 42

600

400

200

0 0

4

2

6

Temps (s)

Fig. 1.4.34 Extrusion of Ti 6–4 (Forecreu, France). XM12/18 1120 12 525 mm P 26 P 27 P 32

Teffort (kN)

12,000

10,000

8000

6000

4000 0

2

4

6

8

10

12

Temps (s)

Fig. 1.4.35 Extrusion of 15–5 pH (Forecreu, France).

be added: soap, MoS2, or graphite. This is a very slow process because of the depression as opposed to compression and its associated risk of striction. Speeds range from 3 to 5 m/mn with a reduction of section of 15%–20% (Figs. 1.4.36 and 1.4.37). l

Focus on rolling

Among rolling technologies, there are two to three main roll techniques. Three-roll rolling has been designed for hard-to-forge material because of its complex structure. Titanium and its alloys have a hexagonal closed packed metallurgical structure, thus a three-roll technique (Fig. 1.4.38) is more efficient (Figs. 1.4.39 and 1.4.40).

Titanium and titanium alloys

85

Fig. 1.4.36 Camera capture (Forecreu, France).

Fig. 1.4.37 Drawing (Forecreu, France).

Fig. 1.4.38 Triangle rolling. (A) Rolled by three roll stand (KOCKS) – first pass. (B) Rolled by three roll stand (KOCKS) – second pass. (C) Rolled by duo stand – second pass.

Fig. 1.4.39 Duo rolling. Rolled by duo stand – first pass.

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Titanium in Medical and Dental Applications

Fig. 1.4.40 Kocks hot rolling (Forecreu, France).

25 mm

13 mm / 12.7 mm

OD

5.5

5.6

6

7

8

9

10

11

12

13

14

15

16

17

18

Fig. 1.4.41 The top values are input material, coil, or bar, bottom is output (Forecreu, France).

The extruded bar is then rolled down using successively triangular reductions of sections with a final rounding, as shown in Figs. 1.4.41 and 1.4.42. l

Focus on wire extraction

The core, a round originally between 10 and 40 mm, has become smaller through extrusion, rolling, and drawing. Differential metallurgical conditions allow the pull out of the wire, usually as small as 0.6–5 mm in bar length form 3–4 m (Fig. 1.4.43). l

Cleaning and dimensional controls

Cannulated bars must be freed from any contamination while final dimensional and cleanliness controls must be conducted. A combination of image recognition for the positioning of the hole and ultrasonic or eddy current NDT as well as fiberscope inspection are conducted (Fig. 1.4.44).

Titanium and titanium alloys

87

Fig. 1.4.42 Rolling mill with eight rolling stands (Forecreu, France).

Fig. 1.4.43 Wire extraction (Forecreu, France).

e

Image capture (NOESIS)

Roughness control (MITUTOYO)

Fibroscopic (CESYCO)

BINOCULAR image x15

Fig. 1.4.44 Controls (Forecreu, France).

Eddy current: to check for transverse cracks. Ultrasonic testing possible (Fig. 1.4.45): l

at ID for defect depth control. in depth to A1 at 1.2 mm flat bottom hole. Metallurgy testing

Micrographic ISO and ASTM specifications (Figs. 1.4.46–1.4.48): -

Free of visible micro inclusion ( 200). α + β equiaxed structure, according to ETTC2 A1–A9 reference pictures.

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Titanium in Medical and Dental Applications

Fig. 1.4.45 US machine (Forecreu, France).

Fig. 1.4.46 Ti 6–4 eli, magnification 200 (Forecreu, France).

Fig. 1.4.47 Hydrogen analysis equipment (Forecreu, France).

Titanium and titanium alloys Nom 17050044-1 tron Description

89 Méthode Methode TITANE

Date de I'analyse 05/07/2017 14:27 Masse de I'échantillon 0,2032 g 0,1842 g 0,1923 g

Hydrogène average 36 ppm Écart-type hydrogène 3 ppm

Hydrogène 36 ppm 40 ppm 34 ppm

Commentaires

n= 3 Hydrogène %RSD 8,65

Date de I'analyse 05/07/2017 14:19 05/07/2017 14:22 05/07/2017 14:27

36 ± 3 Courbe hydrogène ± 1s (ppm) n = 3, RSD(%) = 9

3 2 1 0 0

10

20

30

40

50

60

70

80

Fig. 1.4.48 Hydrogen analysis report (Forecreu, France).

l

Free of α case (hole surface and OD). Free of α network and α platelets. Tensile test

Mechanical tests are performed according to ASTM E8/E8M as specified in the reference specification (ASTM and ISO). Tensile strength, yield strength, elongation, and reduction of area are tested for each and every batch in our laboratory with 200 KN testing equipment. The typical test curve follows (Figs. 1.4.49–1.4.52):

Fig. 1.4.49 Symbolic scheme for sample manufacturing (Forecreu, France).

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Titanium in Medical and Dental Applications

S S0 L

L0

Tube before testing

Tube just before breaking

Fig. 1.4.50 Tubes (Forecreu France).

Fig. 1.4.51 Tensile test equipment (Forecreu, France). 1200 1000

MPa

800 600 400 200 0 0

2

4

6

8

10

12

14

%

N° de l'éprouvette

11

Fm

Rm

Modele

Rp0,2

A 4D

A SD

2%

mm

mm

mm2

KN

MPa

GPa

MPa

%

%

%

2,470

10,000

58,43

60,00

1026,9

103,62

847,2

11,5

35,2

Epaìsseur Dìamètre Section

Pos cossure

Comment :

Entre repères

Fig. 1.4.52 Typical UTS, YS,A% curve Ti 6–4 eli ASTM F136, ISO 5832-3 (Forecreu, France).

Titanium and titanium alloys

91

Fig. 1.4.53 Cannulated bars (Forecreu, France).

1.4.4

Summary

This topic covered many aspects of the production of titanium bars and the processes that follow, including casting, forging, and additive manufacturing to machining technologies. A very special application for fracture repair calls for the use of cannulated bars. The mastering of the entire process requires metallurgical expertise as the composite metallic from billet to final bar permanently uses diverse features of alloy properties. Several forging techniques are being used. Thanks to glass lubrication, extrusion with axial and hydrostatic compression can take place, followed by a triangular axial hot rolling and then an axial traction with drawing (Fig. 1.4.53).

References [1] FORECREU: internal use, www.forecreu.com. [2] PocketBook Orthomaterials, Forecreu, 2012. www.orthomaterials.com.

Transition of surface modification of titanium for medical and dental use

2.1

T. Hanawa Tokyo Medical and Dental University, Tokyo, Japan

2.1.1

Clinical demands and purpose of surface modification

2.1.1.1 Purpose of surface modification meeting clinical demands Many medical devices consisting of metals have been replaced by ceramic and polymers because their properties have been much improved by technological evolution during the last four decades. However, more than 70% of implant devices—including about 95% of orthopedic devices—still consist of metals due to their high strength, toughness, and durability. In particular, titanium and its alloys are successfully used as medical and dental implant materials because of their excellent corrosion resistance and good tissue compatibility. On the other hand, a disadvantage of metals for implant devices is that they are typically artificial materials, so they generally show less biocompatibility and have no biofunction. To improve biocompatibility and add biofunction to metals, surface treatment or surface modification is necessary because biocompatibility and biofunction cannot be added during the manufacturing processes of metals such as melting, casting, forging, and heat treatment. The term “surface modification” is used containing “surface treatment, while surface modification is one category in surface treatment in the field of surface engineering. Surface modification changes the surface composition, structure, and morphology of materials, retaining the bulk mechanical properties intact as shown in Fig. 2.1.1. With surface modification, the biocompatibility of the surface layer can be improved and biofunction can be added. The purpose of surface modification, target devices or arts, and merits are summarized in Table 2.1.1.

2.1.1.2 Bone formation and bone bonding It is well known that titanium has hard-tissue compatibility or bone conduction properties. Hard-tissue compatibility appears based on the interfacial reaction between titanium and hard tissue. In this regards, “osseointegration” is the first definition of the interface phenomenon between titanium and living tissue. The definition of osseointegration is, “The formation of a direct interface between an implant and bone Titanium in Medical and Dental Applications. https://doi.org/10.1016/B978-0-12-812456-7.00005-6 © 2018 Elsevier Inc. All rights reserved.

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Fig. 2.1.1 Surface modification to add target properties to a material retaining the merits of the material.

Properties required to surface of materials Bone formation; Bone bonding Inhibition of bone formation Adhesion to soft tissue Blood compatibility Inhibition of biofilm formation

Surface modification Strength Toughness Durability Elasticity

Properties required to material itself

Table 2.1.1

Requests to metals for medical devices

Required property

Target medical devices

Effect

Bone formation Bone bonding Prevention of bone formation Soft-tissue adhesion

Stem and cup of artificial hip joint; dental implant Bone screw; bone nail

Fixation of devices in bone

Dental implant; trans skin device; external fixation; pacemaker housing All implant devices; treatment tools and apparatus Devices contacting blood

Fixation in soft tissue; prevention of inflectional disease Prevention of infectious disease Prevention of thrombus

Artificial joint

Prevention of generation of wear debris; improvement of durability

Inhibition of biofilm formation Prevention of thrombus Wear resistance

Prevention of assimilation

Coloring

without intervening soft tissue.” No scar tissue, cartilage, or ligament fibers are present between the bone and the implant surface. The direct contact of bone and implant surface can be verified microscopically [1]. This “osseointegration” concept was naturally accepted by dental clinicians and dental materials researchers to explain the biocompatible advantage of titanium among metals. Eventually, titanium occupied a major position in dental implants. After the concept of osseoinegration percolated, studies to elucidate the osseointegration mechanism, especially to investigate the microinterface structure between titanium and bone tissue, have been actively conducted.

Transition of surface modification of titanium for medical and dental use

Top structure Abutment Soft tissue adhesion Fixture Bine bonding

Artificial hip joint Top stem Bone bonding

Head and socket Wear resistance

Dental implant

Bone nail

Liner Bone bonding

97

Internal fixator bone antibonding

Fig. 2.1.2 Medical devices consisting of titanium and its alloys and surface properties required from clinical demands. Bone tissue

Bone tissue

Material Material Chemical bonding between bone and material

Bonding by ingrowth of bone into porous material surface

Chemical adhesion

Mechanical anchoring

Fig. 2.1.3 Bonding manners between bone tissue and material.

Titanium and its alloys are used for dental implants, artificial hip joints, and bone fixators as shown in Fig. 2.1.2. However, their bone formation ability is lower than that of bioactive ceramics such as hydroxyapatite and bioactive glasses. Therefore, numerous studies on surface modification techniques to improve the hard-tissue compatibility of titanium and titanium alloys have been conducted, and some have been commercialized. In the stem of artificial hip joints and dental implants, the chemical bonding of metal surface with bone tissue could not be expected, as shown in Fig. 2.1.3. In other words, it is impossible for metals as typical artificial materials to chemically and naturally bond with bone as living tissue, especially in the human body with bodily fluids. Therefore, surface modification is necessary to accelerate the bone formation and bone bonding on titanium and its alloys.

2.1.1.3 Prevention of bone formation When a titanium alloy is surgically implanted into the human bone, a calcium phosphate layer spontaneously forms on its surface. This ability of a titanium alloy to form calcium phosphate is one of the reasons for its better hard-tissue compatibility than those of other metals; it also accelerates bone formation around itself in the human bone. However, this ability can cause critical complications when the alloys are

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Titanium in Medical and Dental Applications

removed from the bone. When titanium alloys are used for bone fixators such as bone screws and intramedullary nails, titanium alloys form new bone on them and sometimes assimilate with the original bone (osseointegration) [2–4]. Therefore, the bone may be refractured when the fixators are retrieved after bone healing [5–9]. Patients suffering a refracture must be retreated with some difficulty that did not exist at the first treatment. Although stainless steel, which is not equal to titanium alloys in corrosion resistance and strength, may have been used to avoid this problem, refracture during the removal of stainless steel rods also has been reported [10]. Therefore, surface treatments that do not cause bone formation are necessary for the safe utilization of titanium alloys. Surface treatment techniques to inhibit bone formation around titanium alloys have rarely been studied.

2.1.1.4 Soft-tissue adhesion Soft-tissue adhesion is an important property in preventing inflectional disease. “Implantitis” due to inflammation with bacterial invasion is widely known as the cause of failure in dental implants [11]. Bacterial invasion could be prevented with complete adhesion of the junctional epithelium to the abutment part of an implant body (Fig. 2.1.4). In the case of an implant anchor in orthodontics, implant screws pass through the gingiva. If the gingiva does not adhere to the implant anchor, bacteria may invade the interface between the gnathic bone and the implant, inducing inflammation. Also in the external fixator in orthopedics, screws pass through skin. Thus the skin should adhere to the screw. A percutaneous device is also required to adhere to skin. These devices consist of titanium and its alloys. Poor adhesion of soft tissue to titanium and its alloy may induce the invasion of bacteria, inflammation, and eventually inflectional disease. Therefore, surface treatments of titanium and its alloys adhering to soft tissue are required. Bacterial invasion Gingival epithelium

Gingival epithelium

Dental implant

Junctional epithelium Cementum

Alveolar bone

Periodontal ligament

Periodontal tissues around natural tooth

Alveolar bone

Periodontal tissues around implant

Fig. 2.1.4 Periodontal tissues around natural tooth and a dental implant.

Transition of surface modification of titanium for medical and dental use

Surface oxide film (passive film) Mostly amorphous TiO2 with small amount of Ti2O3 and TiO Containing hydroxide and H2O

2-6 nm

99

Fig. 2.1.5 Composition of surface oxide (passive) film on titanium.

Ti substrate

2.1.1.5 Prevention of biofilm formation In the dental implant system, bacteria may invade the postoperative implanted sites (Fig. 2.1.5), generating infections such as peri-implantitis [12–14]. These infections are a serious issue not only in dental implantation but also in orthopedic surgery such as spinal instrumentation [15]. Resistant pathogens such as methicillin-resistant Staphylococcus aureus (S. aureus) cause the infections. The infection is particularly high in patients with a hip prosthesis or implanted spinal or vascular devices [16,17]. Biofilm forms through a complicated process due to multiple factors; initial bacterial adhesion to the surface followed by intercommunication among bacteria leading to the formation of a multilayer of bacteria [18]. Once a biofilm is formed, it is very difficult to remove the glucan-embedded pathogens due to protection by disinfectants. Therefore, surface inhibition of the adhesion of bacteria and help in the removal of bacteria are required. Inhibition of protein adsorption is the most effective technique to inhibit biofilm formation because protein adsorption is usually the initial process of biofilm formation on implant surfaces.

2.1.1.6 Prevention of thrombus The demand of metals for blood-contacting devices is always high. These metals are usually used for long-term intravascular devices such as blood pumps, heart valves, and pacemaker leads as well as for temporary intravascular devices such as guide wires and catheters because of their superior mechanical properties. In metals for intravascular devices, blood compatibility is also important. Adsorption of protein and adhesion of platelets to a metal are important reactions to forming thrombus on the metal and to controlling the blood compatibility of the metal. The adsorption of proteins to a metal surface occurs immediately after the metal contacts the blood, and then platelets adhere to the surface. Because nonprecious conventional metals are usually covered by surface oxide films, the oxide films play an important role against the adsorption of protein and the adhesion of platelets to them. On the other hand, the relation between those and the adhesion behavior of platelets on the metals remains unclear. The level of platelet adhesion on each metal after immersion in a platelet-rich plasma solution was summarized in this order: stainless steel
Co-Cr-Mo alloy < Ti-6Al-4V alloy < Ti-6Al-7Nb alloy < Ti-Ni alloy ¼ Ti [19]. Platelet adhesion

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Titanium in Medical and Dental Applications

is inhibited on stainless steel and the Co-Cr-Mo alloy covered by a Cr2O3-containing passive surface oxide film while accelerated on titanium and its alloys covered by a TiO2-contanining film. A Cr2O3-containing oxide film has smaller relative permittivity than a TiO2-contanining film; it has a larger electrostatic force than the latter; adsorbs more albumins, which work as inhibitory proteins; and inhibits platelet aggregation. Therefore, platelet adhesion and aggregation on titanium and titanium alloys should be controlled by surface modification to inhibit the adsorption of proteins.

2.1.1.7 Increase of wear resistance Friction wear is the most serious problem of sliding parts in medical devices, or artificial joints. The most popular combination of sliding parts consists of a cobalt-chromiummolybdenum (Co-Cr-Mo) alloy and ultrahigh molecular weight polyethylene. Severe wear introduces the obstruction of smooth sliding in the joints. In addition, wear debris induces osteolysis of the femur. In order to decrease the metal wear debris generated from metals that is a predominant problem, the surface must be hardened. The wear resistance of titanium alloys is low, so they cannot be used for bone head applications. The improvement of wear resistance has been attempted through coatings of hard layers such as TiN and DLC (diamond-like carbon). Also, nitrogen-ion implantation to the titanium alloy was attempted. However, these techniques are still not commercialized. At present, titanium and its alloys with enough wear resistance are not available, so Co-Cr-Mo alloys are used as sliding parts.

2.1.1.8 Coloring Coloring of titanium and its alloys is carried out through anodizing. Coloring is often utilized in the surface finishing of products, especially for medical devices. Surface coloring of the components of the devices allows the easy identification of devices and parts. This benefit is great in surgical applications. Anodizing is a process that controls and adjusts the oxide thickness of metal surfaces. This adjustment changes the spectrum of light, resulting in perceived color. By precisely controlling the surface oxide thickness, an entire range of colors could be produced. In medical and dental devices, the coding of color for size allows instant recognition of desired parts when time and accuracy are critical. Implant devices, medical instruments, and device components can be coded with standardized or specialty colors to increase accuracy and efficiency during surgery.

2.1.2

Surface of titanium

2.1.2.1 Passive film With the exception of reduction environments, a reaction film on metals is always formed by the corrosion process. Passive film is such a reaction film, and it is particularly significant for corrosion protection. When solubility of the film is extremely low

Transition of surface modification of titanium for medical and dental use

101

and pores are absent, film adhesion to the substrate will be strong. The film then becomes corrosion resistant or a passive film. A passive film is a few nm thick and transparent. Because the formation rate is tremendously fast, a passive film readily becomes amorphous. For example, an oxide film on a titanium metal substrate was generated in 30 ms [20]. Due to the absence of a grain boundary and structural defects in amorphous films, they are corrosion-resistant. However, corrosion resistance of the film decreases if the crystallinity of the film is increased. Fortunately, passive films contain water molecules that promote and maintain amorphousness. When titanium is polished in de-ionized water and analyzed immediately using X-ray photoelectron spectroscopy, the Ti 2p electron energy region peak obtained from the titanium gives four doublets originating from the valences Ti0, Ti2+, Ti3+, and Ti4+, respectively. The surface oxide contains hydroxide or hydroxyl group (OH) and water [21,22]. The surface film on titanium consists mainly of amorphous or low-crystalline and nonstoichiometric TiO2 (Fig. 2.1.5), and the film is protective against chloride ions.

2.1.2.2 Surface hydroxyl groups The surface of oxide reacts with moisture in the air and hydroxyl groups are rapidly formed. In the case of titanium, the surface oxide immediately reacts not only with water molecules in aqueous solutions but also with moisture in the air and is covered by hydroxyl groups [23,24] as shown in Fig. 2.1.6A. The surface hydroxyl groups contain both terminal OH and bridge OH in equal amounts. Active surface hydroxyl groups on the oxide film dissociate in aqueous solutions and form electric charges as shown in Fig. 2.1.6B [23–25]. A positive or negative charge due to dissociation is governed by the pH value of the surrounding aqueous

Terminal OH

H

H

H

O •• + d−

O

d+

Ti

H O ••

d−

O

O

Ti

Bridge OH

H d+ •• d −

O

O

O

H

Ti

O

(A) OH− M+

(B)

Smaller pH

H+ OH

O−

M

M

Point of zero charge Pzc

Larger pH

Fig. 2.1.6 Formation process of hydroxyl group on titanium oxide (A) and dissociation of the hydroxyl group in an aqueous solution and point of zero charge (pzc) (B).

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Titanium in Medical and Dental Applications

solution: positive and negative charges are balanced at a certain pH where the apparent charge is zero. This pH is the point of zero charge (pzc). The pzc is the unique value for an oxide and an indicator, which the oxide surface shows as an acidic or basic property. For example, in the case of TiO2, the pzc of rutile is 5.3 and that of anatase is 6.2 [23]. In other words, the anatase surface is acidic at a smaller pH and basic at a larger pH than 6.2. Active surface hydroxyl groups and electric charges formed play important roles for the immobilization of molecules. Therefore, the concentration of the surface hydroxyl group on the oxide and pH of the surrounding solution is an important factor for the immobilization of molecules.

2.1.2.3 Calcium phosphate formation on titanium The composition of surface oxide film varies according to the change in environment while the film is macroscopically stable. Passive surfaces coexist in close contact with electrolytes and partial dissolution and reprecipitation of the surface continuously occur from the microscopic viewpoint. In this sense, surface composition is always changing according to the environment. Calcium phosphates form on titanium and its alloys by immersion in Hanks’ solution and other solutions [26–29]. This phenomenon is also observed under cell culture [29]. The above phenomena are characteristic of titanium and its alloys [27]. The characterization of dental implants retrieved from a human jawbone also reveals the formation of calcium phosphate as well as sulfur on the surfaces [30,31]. The surface oxide film on titanium is not completely oxidized and is slightly reactive while that on zirconium is relatively stable. The film on zirconium is also more passive and protective than that on titanium. Neither calcium nor phosphate stably exists alone on titanium; stable and protective calcium phosphate is formed on titanium in biological environments. On the other hand, calcium is never incorporated on zirconium and zirconium phosphate is formed. The zirconium phosphate formed on zirconium is highly stable and establishes a protective layer; therefore, no calcium reacts with the layer [32]. The formation capabilities of calcium phosphate in titanium, zirconium, tantalum, and niobium are evaluated [33].

2.1.2.4 Protein adsorption matter to titanium When a metal is implanted into the human body and contacts living tissue, proteins are instantly adsorbed to the metal surface. Adsorption of proteins influences the adhesion of cells to the surface. Denaturalization and fragmentation of adsorbed protein may affect the function of the host body (Fig. 2.1.7). To characterize proteins adsorbed to metals and metal oxides, various techniques can be used [34], especially that of ellipsometry [35]. To predict protein adsorption, the measurement of wettability is used where a liquid droplet is applied to the metal. Fibrinogen is much more naturally adsorbed on a titanium surface than on a gold surface [36] because the dielectric

Transition of surface modification of titanium for medical and dental use

103

Unadsorbed protein Body fluid Adsorbed protein Polar group

Surface oxide film

Fig. 2.1.7 Possible model of protein adsorption by surface oxide film on titanium.

constant, the factor governing electrostatic force, of TiO2 is 80.1 and similar to that of water. Therefore, fibrinogen retains its conformation even after the adsorption on a titanium surface.

2.1.2.5 Mechanism of hard tissue compatibility in titanium Titanium and its alloys are well known as being among the best biocompatible materials, and they are successfully used for orthopedic and dental implants. Why do titanium and its alloys show such good biocompatibility compared to other alloys? The explanation is generally believed to be that titanium passivates in air and aqueous solutions and that passive film is stable even in the human body. Therefore, it was first demonstrated that the good hard-tissue compatibility of titanium is caused by its high corrosion resistance. It is known now that this hypothesis is false. Electroplating of platinum on titanium makes delay bone formation on itself, while the corrosion resistance increased [37]. Therefore, the good hard-tissue compatibility of titanium is caused not only by its high corrosion resistance but also by other factors. As described above, the composition of the surface oxide film varies according to the change in the environment in spite of the macroscopic stability. The composition and properties of the oxide film regenerated in a biological environment are different from those in water [38]. As described above, titanium naturally forms calcium phosphate on itself. Therefore, bone formation is faster on titanium implanted in hard tissue simply because the surface oxide film is titanium oxide. Protein adsorption influences cells adhesion. The titanium surface is almost medium and does not show an extremely positive or negative charge. Therefore, the influence of proteins that adsorb to titanium is smaller than those to other metals. In addition, protein adsorbs with little conformational change to the titanium surface, so the property of the protein remains after adsorption. As described above, many researchers have attempted to elucidate the mechanism of osseointegration by characterization of titanium surface oxide (composition and change of it), surface hydroxyl groups, adsorption of proteins (amount, speed, change in the conformation, and denaturalization), and adhesion, proliferation, and differentiation of cells. However, the true mechanism of osseointegration is still not clear.

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Titanium in Medical and Dental Applications

2.1.3

Surface modification techniques

2.1.3.1 Overview To improve biocompatibility and add biofunction, surface modification is essential, as explained previously. On the other hand, porous or roughened surfaces are usually effective to bind bone tissue to titanium materials. Human tissue, such as bone, is expected to grow into the rough and porous surface; finally the materials and bone are strongly connected as a result of the so-called anchoring effect. Fig. 2.1.3 shows chemical bonding and a mechanical anchoring connection between bone and material. The current surface modification techniques to improve the hard-tissue compatibility focus on the effect of chemical bonding and/or mechanical anchoring to the metal substrate. The surface modification techniques studied in the field of biomaterials are summarized in Fig. 2.1.8. Some of them are commercialized already. Surface modification techniques are reviewed elsewhere [39–41].

2.1.3.2 Category of surface modification According to the purpose of surface modification (see Table 2.1.1), the most suitable modification technique should be selected. Surface modification techniques are categorized as follows. HA

HA Plasma spray

TiO2

Electrochemical coating

TiO2

Blast MAO Porous layer

Acid etching

HA TiO2

PVD

Dry process

Chemical treatment

H2O2 immersion Alkaline immersion

CaTiO3

Wet process

Sol-Gel

+ Heating

+ Hydrothermal

HA CVD Electrochemical modification

DLC

Cathodic treatment Thermo-electrochemical

N ion He ion

lon implantation

Immersion of biomolecules and biofunctional molecules

PEG

Polymer coating

Gelatin, hydrogel

Ca ion Noble metal ion

Peptide, collagen collagen

Fig. 2.1.8 Surface modification techniques studied in the field of biomaterials; some of them are commercialized.

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2.1.3.2.1 Dry process and wet process Dry processes (performed in air, decreased pressure, and vacuum) and wet processes (performed in aqueous solutions) are conventional. Blast and bead sintering are mechanical processes to form rough surfaces. The dry process can be categorized according to the effects on solid surface: film formation, sputtering, and ion implantation. When an ion impacts a solid surface, attaching, sputtering, and implantation effects occur on the surface according to the ion’s energy. Spray is a useful technique to form a rough TiO2 layer and hydroxyapatite (HA) layers. These techniques are successfully commercialized. Also, PVD and CVD are useful techniques to form thin and smooth layers. On the other hand, ion implantation is used for the formation of a thermal-unequilibrium layer. This surface-modified layer is sometimes effective in accelerating bone formation [42]. The disadvantage of this process is that instruments are expensive. Therefore, the dry process is suitable for mass production. The wet process is carried out in aqueous solutions. This process does not require large facilities or high costs. In the case of the wet process, the modifications are performed in aqueous solutions with immersion and electrochemical processes. The wet process does not require special instruments and is easy to conduct in a laboratory level. The resultant modified layer changes according to changes in the following parameters: l

l

l

Composition and pH of aqueous solution. Potential gain due to electrolysis. Current density of electrolysis.

In acid etching, a rough surface can be obtained. Immersion in an alkaline solution is effective to accelerate bone formation; this technique is commercialized [43]. Electrochemical techniques such as anodic oxidation, cathodic polarization, and immobilization of biofunctional molecules are described later.

2.1.3.2.2 Surface layer From the viewpoint of the resultant thickness and composition of the surface layer, surface modification techniques are categorized into three types: Thin-layer coating, surface-modified layer formation, and immobilization of biofunctional molecules (Fig. 2.1.9). Thin-layer coating is a typical and a conventional surface modification technique. Thin and uniform composition layers usually with a few micrometers thickness are formed by plasma spray, PVD, CVD, anodic oxidation, etc. Surfacemodified layer formation is performed by ion implantation, alkaline treatment [43], and cathodic polarization. Immobilization of biofunctional molecules is performed by chemical treatment and electrodeposition.

2.1.3.2.3 Calcium phosphate formation To accelerate bone formation, the most popular technique is the coating of calcium phosphate such as HA. Plasma spraying of apatite on metals is widely used to form the apatite layer that is the nucleus for active bone formation and conductivity.

106

Fig. 2.1.9 Category according to the resultant thickness and composition of the surface layer: Thin layer coating, surface-modified layer formation, and immobilization of biofunctional molecule.

Titanium in Medical and Dental Applications

Thin layer coating

Surface modified layer formation

Immobilization of biofunctional molecules

However, in the case of plasma-sprayed HA, the HA-titanium interface or HA itself may fracture under relatively low stress because of low bonding strength and low toughness of the sprayed HA layer [44]. The other techniques to improve toughness and bonding of the HA layer have been attempted. On the other hand, hard-tissue compatibility could be improved by the modification of the titanium surface instead of coating by the HA layer. In this regard, various techniques to modify the titanium surface are given as follows: l

l

l

l

Immersion in alkaline solution and heating. Immersion in hydrogen peroxide solution. Immersion and hydrothermal treatment in calcium-containing solution. Calcium ion implantation.

The most famous technique is alkaline treatment. By immersion of titanium into NaOH or KOH alkaline solutions, a hydrated titanium oxide gel containing alkaline ions with 1-μm thickness is formed on the titanium substrate [43]. After heating, the gel layer condenses and strongly bonds to the substrate.

2.1.3.2.4 Chemical bonding and anchoring The current surface modification techniques to improve hard-tissue compatibility focus on the strengthening of chemical bonding and/or the increase of mechanical anchoring as explained already. Surface roughness usually changes by the surface modification process. Even targeting chemical bonding, surface morphology always changes; there is sometimes a mechanical anchoring factor (Fig. 2.1.10). Therefore, it is difficult to distinguish the chemical effect and the anchoring effect when surface modification is performed.

2.1.3.2.5 Cell adhesion According to the purpose of surface modification, cell adhesion is required or is not required as shown in Fig. 2.1.11. To accelerate bone formation and bone bonding and soft tissue adhesion, cells must adhere stably, differentiate, and generate tissues rapidly. On the other hand, to inhibit bacterial adhesion, the formation of biofilm, platelet adhesion, and the formation of thrombus, cell adhesion must be prevented. Therefore, these properties are completely opposite from the viewpoint of cell adhesion.

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Surface modification to improve chemical composition

Surface modification to increase roughness Blast Acid etching Anodic oxidation etc.

HA coating Alkaline treatment etc.

Comparing with a control, cell proliferation, adhesion, calcification, bone bonding, etc. are enhanced

Effect of not only chemical composition but also roughness

Change in roughness

Fig. 2.1.10 Change in surface roughness even with surface modification to improve the surface composition.

Bacteria

Cell

Protein Small molecule Tissue Metal

Fig. 2.1.11 Cellular and bacterial adhesion to the metal surface after protein adsorption.

Surface oxide Metal substrate Time

2.1.3.3 Electrodeposition and electrochemical techniques Electrodeposition is categorized into the following three processes: (1) electroplating, (2) electrophoretic deposition, and (3) underpotential deposition. For biomedical materials, ceramics and polymers are electrodeposited on metallic materials. The target substrates for electrodeposition are titanium and its alloys containing nickeltitanium superelastic and shape memory alloy; the electrodeposited coating materials are ceramics such as HA, octacalcium phosphate (OCP), brushite (DCPD), other calcium phosphates (Ca-P), DLC, and TiO2, metals such as magnesium and tantalum, and polymers such as collagen, chitosan, and poly(ethylene glycol) (PEG). The combinations between target substrates and electrodeposited coating materials and the purposes in electrodeposition are summarized in Table 2.1.2. These combinations change according to the mission of the resultant materials. HA coating is usually performed for the improvement of hard-tissue compatibility (bone formation and bone bonding). DLC coating is performed for the improvement of friction wear. TiO2 coating is performed for the improvement of corrosion resistance and hard-tissue compatibility, especially bone bonding by the forming of connective porous layers. A thick TiO2 layer with connective pores is produced by microarc oxidation (MAO).

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Table 2.1.2 Electrodeposited layers to titanium and it alloys and the purpose Deposited material

Purpose

Hydroxyapatite Calcium phosphate Calcium compound TiO2 nanotube Carbon nanotube Diamond-like carbone

Bone formation Bone bonding Bone formation Bone bonding Friction coefficient Wear resistance Bone formation Bone bonding Bone formation Corrosion resistance Antibacterial property Antibacterial property Inhibition of platelet adhesion Bone formation

TiO2 and ZrO2 by micro-arc oxidation Biomolecule

Poly(ethylene grycol)

Biomolecule and biofunctional molecules are immobilized to a metal surface to add biofunctions to the metal. Electrodeposition in titanium and its alloys is summarized in Fig. 2.1.12. Studies on the electrodeposition process of biomedical materials are reviewed elsewhere [45].

2.1.3.4 Immobilization of biofunctional molecules The immobilization of biofunctional molecules to metal surfaces is effective to add biofunctions to the metal. The hard-tissue compatibility and soft-tissue compatibility of metals are sometimes required where metals are used for parts of implants. Stents are placed in stenotic blood vessels for dilatation, and blood compatibility is Fig. 2.1.12 Graphic summary of the coating layer formed by electrodeposition.

Anodic

Cathodic Pulsed Electrodeposition

Ca, P, Ag, Mg, Sr, Zn, Cu, Ag, Si, Pt Doped TiO2

Coating layer HA

DCPD

HA/OCP

Ca-P

TiO2 nanotube Ti

Ti-6AI-4V

Ti-Nb-Zr

Ti-6AI-7Nb

Ti–29Nb–13Ta–4.6Zr Ti-24Nb-4Zr-7.9Sn

Substrate

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109

necessary. Furthermore, guide wires and catheters require lubrication in blood vessels for proper sliding. When metals are used as sensing devices, the adhesion of cells must be controlled. The major cause of the retrieval of implants during the service is infection due to biofilm formation. Therefore, a biofilm-inhibiting surface is required, the fundamental property of which is to control the adsorption of proteins and adhesion of cells, platelets, and bacteria. This functional surface is possibly created by the immobilization of biofunctional molecules. This technique makes it possible to apply metals to a scaffold in regenerative medicine. Some metal-polymer composites are reviewed in other textbooks [46–48].

2.1.4

Transient of surface modification

By referring to the stem of artificial hip joints and the fixture of dental implants, time transients of surface modification techniques to improve bone formation and bone bonding at the research level and commercial level are shown in Fig. 2.1.13. l

l

l

l

l

First generation: Machining surface. Second generation: Morphological surface. Third generation: Physicochemical active surface. Fourth generation: Biochemical active surface. Fifth generation: Biological surface?

Surface modifications in the fourth generation were easily accepted by researchers because immobilization of biomolecules accelerating bone formation is easily vel

ial le

ress

Prog

erc omm

5th Generation?

on c

4th Generation 3rd Generation

Biochemical active surface Physicochemical Immobilization of collagen, BMP, active surface peptide, gelatin HA coating etc. alkaline treatment etc.

2nd Generation 1st Generation

Morphological surface Blast Mechanical surface Groove Gliding Etching Anodic oxidation etc. ress

Prog

Biological surface Coating of tissue, Stem cell etc.

vel

ch le

ear n res

o

Fig. 2.1.13 Time transient of surface modification techniques to accelerate bone formation and estrangement in surface modification techniques between research and commercialization.

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understood. Therefore, many studies have been attempted. However, to commercialize the immobilization of biofunctional molecules, it is necessary to credit the safety, maintenance of quality during storage, and dry-conditioned durability of the immobilized layer. It is difficult for manufacturers to commercialize unless they find the value of commercialization. There are many problems to commercialize the immobilized materials while it is easy to show good results in the basic research. Eventually, most commercialized goods are categorized into the second generation. There are a few commercialized goods of the third generation, but there is no prospect for commercialization in the fourth generation, at present. This mismatch between research and commercialization is not always caused by the delay of permission for commercialization. Another reason for the delay of the commercialization of the third generation goods is possibly because goods employing mechanical anchoring with the rough surface expecting bone ingrowth to the porous surface is practically better than goods employing chemical bonding with bone.

2.1.5

Application to regenerative medicine

Titanium is widely used in medical implants and their surfaces may be biofunctionalized through various techniques such as dry and wet processes including immobilization of biomolecules. The major purpose of surface modification is the improvement of hard-tissue compatibility or the acceleration of bone formation. On the other hand, the electrodeposition is useful for all electroconductive and morphological materials not only to inhibit platelet adhesion and bacterial adhesion but also to enhance bone formation. These techniques make it possible to apply metals to a scaffold in regenerative medicine. Probably, fourth-generation techniques in Fig. 2.1.13 may be used to perform surface modification of the fifth generation. For scaffolds, titanium fiber, foil, and mesh are useful.

2.1.6

Future of surface modification

Metals are widely and safety used in medicine not only for orthopedic implants but also for cardiovascular devices and other purposes. Implant devices are always used in contact with living tissues. Therefore, interactions between material surfaces and living tissues must be well understood. This knowledge is mandatory to develop new novel materials. In particular, reactions between biomolecules or cells and metal surfaces are important. A good knowledge base of these reactions always helps to add biofunctions to metals. If both bone formation and antibacterial properties or both soft-tissue adhesion and antibacterial properties are acquired simultaneously, the material is an optimal and frontier multifunctional material. In addition, a sensing function may be added to the titanium surface by a surface modification technique.

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References [1] P.I. Br€anemark, P.I.B.O. Hansson, R. Adell, U. Breine, J. Lindstr€ om, O. Halle˘n, A. Ohman, Osseointegrated implants in the treatment of the edentulous jaw. Experience from a 10-year period, Scand. J. Plast Reconstr. Surg. Suppl. 16 (1977) 1–132. [2] T. Jinno, V.M. Goldberg, D. Davy, S. Stevenson, Osseointegration of surface-blasted implants made of titanium alloy and cobalt-chromium alloy in a rabbit intramedullary model, J. Biomed. Mater. Res. 42 (1998) 20–29. [3] J.E. Feighan, V.M. Goldberg, D. Davy, J.A. Parr, S. Stevenson, The influence of surfaceblasting on the incorporation of titanium-alloy implants in a rabbit intramedullary model, J. Bone Joint Surg. Am. 77 (1995) 1380–1395. [4] T. Jinno, S.K. Kirk, S. Morita, V.M. Goldberg, Effects of calcium ion implantation on osseointegration of surface-blasted, J. Arthroplast. 19 (2004) 102–109. [5] M. Takakuwa, M. Funakoshi, K. Ishizaki, T. Aono, H. Hamaguchi, Fracture on removal of the ACE tibial nail, J. Bone Joint Surg. (Br.) 79 (1997) 444–445. [6] D.H. Jones, G. Schmeling, Tibial fracture during removal of a tibial intramedullary nail? J. Orthop. Trauma 13 (1999) 271–273. [7] C.J. Seebauer, K.M. van Scherpenzeel, N.P. Haas, H.J. Bail, Tibia fracture following removal of the ETN (expert tibia nail): a case report, Arch. Orthop. Trauma Surg. 129 (2009) 949–953. [8] P.L. Sanderson, W. Ryan, P.G. Turner, Complications of metalwork removal, Injury 23 (1992) 29–30. [9] H. Young, T. Claire, Complications associated with the use of a titanium tibial nail, Injury 38 (2007) 223–226. [10] M. Aizawa, T. Jinno, N. Nanke, K. Ishii, S. Kawachi, Femoral fracture caused by removal of femoral intramedullary nail made of stainless steel, J. Med. Case 6 (2015) 105–108. [11] P.J. Pe`rez-Chaparro, P.M. Duarte, J.A. Shibli, S. Montenegro, S.L. Heluy, L.C. Figueiredo, M. Faveri, M. Feres, The current weight of evidence of the microbiologic profile associated with peri-implantitis: a systematic review, J. Periodontol. 87 (2016) 1295–1304. [12] S. Dibart, M. Warbington, M.F. Su, Z. Skobe, In vitro evaluation of the implant-abutment bacterial seal: the locking taper system, Int. J. Oral Maxillofac. Implants 20 (2005) 732–737. [13] R. Glauser, P. Schupbach, J. Gottlow, C.H.F. Hammerle, Periimplant soft tissue barrier at experimental one-piece mini-implants with different surface topography in humans: a light-microscopic overview and histometric analysis, Clin. Implant. Dent. Relat. Res. 7 (2005) S44–S51. [14] M. Tesmer, S. Wallet, T. Koutouzis, T. Lundgren, Bacterial colonization of the dental implant fixture-abutment interface: an in vitro study, J. Periodontol. 80 (2009) 1991–1997. [15] E.E. MacKintosh, J.D. Patel, R.E. Marchant, J.M. Anderson, Effects of biomaterial surface chemistry on the adhesion and biofilm formation of Staphylococcus epidermidis in vitro, J. Biomed. Mater. Res. A 78 (2006) 836–842. [16] P. Sendi, M. Rohrbach, P. Graber, R. Frei, P.E. Ochsner, W. Zimmerli, Staphylococcus aureus small colony variants in prosthetic joint infection, Clin. Infect. Dis. 43 (2006) 961–967. [17] J. Barberan, L. Aguilar, M.J. Gimenez, G. Carroquino, J.J. Granizo, J. Prieto, Levofloxacin plus rifampicin conservative treatment of 25 early staphylococcal infections of osteosynthetic devices for rigid internal fixation, Int. J. Antimicrob. Agents 32 (2008) 154–157.

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[18] S.E. Cramton, C. Gerke, N.F. Schnell, W.W. Nichols, F. G€ otz, The intercellular adhesion (ICA) locus is present in Staphylococcus aureus and is required for biofilm formation, Infect. Immun. 67 (1999) 5427–5433. [19] Y. Tanaka, K. Kurashima, H. Saito, A. Nagai, Y. Tsutsumi, H. Doi, N. Nomura, T. Hanawa, In vitro short-term platelet adhesion on various metals, J. Artif. Organs 12 (2009) 182–186. [20] T. Hanawa, Surface modification of metallic biomaterials, in: S.H. Teoh (Ed.), Engineering Materials for Biomedical Applications, World Scientific, Hackensack, NJ, USA, 2004, pp. 4.1–4.36. [21] K. Asami, S.C. Chen, H. Habazaki, K. Hashimoto, The surface characterization of titanium and titanium-nickel alloys in sulfuric acid, Corros. Sci. 35 (1993) 43–49. [22] T. Hanawa, S. Hiromoto, K. Asami, O. Okuno, K. Asaoka, Surface oxide films on Ti alloys regenerated in hanks’ solution, Mater. Trans. 43 (2002) 3000–3004. [23] J. Westall, H. Hohl, A comparison of electrostatic models for the oxide/solution interface, Adv. Colloid Interf. Sci. 12 (1980) 265–294. [24] T.W. Healy, D.W. Fuerstenau, The oxide-water interface-interreaction of the zero point of charge and the heat of immersion, J. Colloid Sci. 20 (1965) 376–386. [25] H.P. Boehm, Acidic and basic properties of hydroxylated metal oxide surfaces, Discuss. Faraday Soc. 52 (1971) 264–289. [26] T. Hanawa, M. Ota, Calcium phosphate naturally formed on titanium in electrolyte solution, Biomaterials 12 (1991) 767–774. [27] T. Hanawa, Titanium and its oxide film: a substrate for formation of apatite, in: J.E. Davies (Ed.), The Bone-Biomaterial Interface, University of Toronto Press, Toronto, 1991, pp. 49–61. [28] T. Hanawa, M. Ota, Characterization of surface film formed on titanium in electrolyte, Appl. Surf. Sci. 55 (1992) 269–276. [29] S. Hiromoto, T. Hanawa, K. Asami, Composition of surface oxide film of titanium with culturing murine fibroblasts L929, Biomaterials 25 (2004) 979–986. [30] J.E. Sundgren, P. Bodo, I. Lundstrom, Auger electron spectroscopic studies of the interface between human tissue and implants of titanium and stainless steel, J. Colloid Interface Sci. 110 (1986) 9–20. [31] M. Espostito, J. Lausmaa, J.M. Hirsch, P. Thomsen, Surface analysis of failed oral titanium implants, J Biomed Mater Res B Appl Biomater 48 (1999) 559–568. [32] Y. Tsutsumi, D. Nishimura, H. Doi, N. Nomura, T. Hanawa, Difference in surface reactions between titanium and zirconium in Hanks’ solution to elucidate mechanism of calcium phosphate formation on titanium using XPS and cathodic polarization, Mater. Sci. Eng. C 29 (2009) 1702–1708. [33] Y. Tsutsumi, D. Nishisaka, H. Doi, M. Ashida, P. Chen, T. Hanawa, Reaction of calcium and phosphate ions with titanium, zirconium, niobium, and tantalum, Surf. Interface Anal. 47 (2015) 1148–1154. [34] B. Ivarsson, I. Lundstr€om, Physical characterization of protein adsorption on metal and metal oxide surfaces, CRC Crit. Rev. Biocompat. 2 (1986) 1–96. [35] H. Elwing, Protein absorption and ellipsometry in biomaterial research, Biomaterials 19 (1998) 397–406. [36] J.E. Sundgren, P. Bod€o, B. Ivarssonm, I. Lundstr€om, Adsorption of fibrinogen on titanium and gold surfaces studied by ESCA and ellipsometry, J. Colloid Interface Sci. 113 (1986) 530–543. [37] Y. Itakura, T. Tajima, S. Ohoke, J. Matsuzawa, H. Sudo, S. Yamamoto, Osteocompatibility of platinum-plated titanium assessed in vitro, Biomaterials 10 (1989) 489–493.

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[38] T. Hanawa, K. Asaoka, K. Asami, Repassivation of titanium and surface oxide film regenerated in simulated bioliquid, J. Biomed. Mater. Res. 40 (1998) 530–538. [39] D.M. Brunette, P. Tengvall, M. Textor, P. Thomse (Eds.), Titanium in Medicine, Springer, Berlin, 2001, pp. 231–455. [40] T. Hanawa, An overview of biofunctionalisation of metals in Japan, J. R. Soc. Interface 6 (2009) S361–S369. [41] T. Hanawa, Biofunctionalization of titanium for dental implant, Jpn. Dent. Sci. Rev. 46 (2010) 93–101. [42] T. Hanawa, Y. Kamiura, S. Yamamoto, T. Kohgo, A. Amemiya, H. Ukai, K. Murakami, K. Asaoka, Early bone formation around calcium-ion-implanted titanium inserted into rat tibia, J. Biomed. Mater. Res. 36 (1977) 131–136. [43] H.M. Kim, F. Miyaji, T. Kokubo, T. Nakamura, Preparation of bioactive Ti and its alloys via simple chemical surface treatment, J. Biomed. Mater. Res. 32 (1996) 409–417. [44] J.L. Ong, L.C. Lucas, Post-deposition heat treatment for ion beam sputter deposited calcium phosphate coatings, Biomaterials 15 (1994) 337–341. [45] T. Hanawa, Electrodeposition of calcium phosphates, oxides, and molecules to achieve biocompatibility of metals. Module Chem. Mol. Sci. Chem. Eng (2016). https://doi.org/ 10.1016/B978-0-12-409547-2.11159-X. [46] W. Possart, Adhesion of polymers, in: J.F. Helsen, H.J. Breme (Eds.), Metals as Biomaterials, Wiley, New York, 1998, pp. 197–218. [47] H. Worch, Special thin organic coatings, in: J.F. Helsen, H.J. Breme (Eds.), Metals as Biomaterials, Wiley, New York, 1998, pp. 177–196. [48] T. Hanawa, Metal-polymer composite biomaterials, in: S. Dumitriu, V. Popa (Eds.), Polymeric Biomaterials, vol. 1, CRC Press, Boca Raton, FL, 2013, pp. 343–375.

Modern techniques of surface geometry modification for the implants based on titanium and its alloys used for improvement of the biomedical characteristics

2.2

E.G. Zemtsova, A.Y. Arbenin, R.Z. Valiev, V.M. Smirnov St. Petersburg State University, St. Petersburg, Russia

2.2.1

Introduction

Bone implantation with the application of titanium implants became possible after the works of Prof. Bra˚nemark in the early 1960s, where titanium osseointegration was established for the first time [1]. After this discovery, it was found that such an effect is characteristic not only for titanium but also for zirconium oxide [2], various alloys [3], hydroxyapatite [4], and other materials. However, despite this, the majority of bone implants are made of titanium and its alloys. This is due to both biomedical and mechanical characteristics as well as commercial availability. More than 50 years of experience suggests the main feature of titanium osseointegration: the geometric structure of the surface is the most important factor influencing the implant engraftment along with the chemical composition. This is demonstrated using the results of both cytological and histological studies [5,6]. In general, the majority of the works describe a positive contribution of the developed surface topography into biomedical properties related to the polished titanium [7–9]. However, in some cases insignificant relief modification causes the valuable response of osteoblasts and bone tissue [10]. We try to disclose these relationships based on the works of our colleagues. Also, in this chapter, we try to analyze the existing techniques of the surface geometry modification. Based on the reviewed materials, we will outline the prospective of the surface geometry modification of titanium implants without changing the chemical composition of the surface layer for the improvement of their biomedical characteristics.

Titanium in Medical and Dental Applications. https://doi.org/10.1016/B978-0-12-812456-7.00006-8 © 2018 Elsevier Inc. All rights reserved.

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2.2.1.1 The effect of surface geometry on the biomedical characteristics of titanium implants The process of osseointegration is highly dependent on surface topography. This effect is due to the specific features of the interaction between living tissue and the implant surface during engraftment. In the work [11], the interaction is described between the titanium surface of various roughness and incubated cells. The investigated sample comprises gradiently polished sandblasted aluminum coated by the titanium layer. As a result, authors fabricated a plate with a roughness gradient. To assess the cellular response, osteoblasts and fibroblasts (i.e., young cells of connective tissue) were incubated on samples (Fig. 2.2.1). Subsequently, an integrated amount of cells on the sample was analyzed. The important conclusion is made based on the following dependence: when the surface roughness increases, the proliferation rate increases for osteoblasts and decreases for fibroblasts (Fig. 2.2.2). Therefore, rough surface titanium is more suitable for bone implantation because it promotes bone tissue growth and protects implants from encapsulation by the connective tissue. These results on the roughness influence were confirmed also in [12–16]. However, the increase of micron-sized roughness is not the only cause of the improvement of biomedical properties. The roughness firstly prevents the growth of fibroblasts and the drift of osteoblasts. The drift occurs on the flat surfaces for a

sb Ra = 5.70 µm

01 mm Ra = 4.50 µm

04 mm Ra = 2.48 µm

09 mm Ra = 1.12 µm

Fig. 2.2.1 Microphotographs of the surface of various roughness before incubation (A) and after 7 days incubation of osteoblasts (B) and fibroblasts (C) [11]. Reprinted from T.P. Kunzler, T. Drobek, M. Schuler, N.D. Spencer, Systematic study of osteoblast and fibroblast response to roughness by means of surface-morphology gradients, Biomaterials 28 (2007) 2175–2182. Copyright (2007), with permission from Elsevier.

* *

Cell number ¥103 [#/cm2]

Cell number ¥103 [#/cm2]

25

6

Day 7 Day 4 Day 2 Day 1

*

20 15 *

*

*

10

*

5

**

4 3 2 **

***

1

**

0

0 S T

(A)

5

Day 7 Day 4 Day 2 Day 1

Modern techniques of surface geometry modification

30

1

2

3 Ra [mm]

4

5

S T

6

(B)

1

2

3

4

5

6

Ra [mm]

Fig. 2.2.2 Rate of proliferation of osteoblasts (A) and fibroblasts (B) on the surface with the roughness gradient [11]. Reprinted from T.P. Kunzler, T. Drobek, M. Schuler, N.D. Spencer, Systematic study of osteoblast and fibroblast response to roughness by means of surface-morphology gradients, Biomaterials 28 (2007) 2175–2182. Copyright (2007), with permission from Elsevier.

117

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long time. For the rough surface, drift time is decreased and this leads to the transformation of the cytoskeleton for better adhesion [17]. Note that there is no straight dependence of the proliferation and roughness on the whole size range. The osteoblasts incubation experiment [18] was performed on the surface of compressed monodisperse titanium oxide powders (Fig. 2.2.3). As one can see, the proliferation reaches maximum on the samples with high micron-sized roughness and developed nanorelief. This indicates various processes during sorption and proliferation of the osteoblasts on micro- and nanorelief. In general, this is predictable because the tissues are not homogeneous by default: there is a molecular level of organization, an extracellular matrix with submicron-sized elements, and cells of micron size. One should not expect the equivalent response on the relief of a different scale. After reviewing cytological responses on the micron-sized relief, let’s discuss the same for nanorelief. A number of articles describe the efficiency of this surface layer organization (e.g., [19,20]), however, there are no attempts to explain this effect. To understand the process, more interesting are the systematic studies carried out with precision-controlled geometry. Applied methods are likely to have no practical prospects in implantology, but they help to understand the mechanisms of the surface processes. Lithography with reactive ionic etching [21] was used for the formation of periodic structures that simulate bundles of collagen. These data indicate the effectiveness of mimicry and precision of adjustment to the real structure of collagen (Fig. 2.2.4). It is also worthy to note that the reference sample

Cell density (cells/cm2)

4000

* 3000 * * 2000

1000

0 Glass

2120

32 97 56 Titania grain size (nm)

26

20

Fig. 2.2.3 The dependence between osteoblasts proliferation and the mean grain size of titanium oxide [18]. Reprinted from T.J. Webster, R.W. Siegel, R. Bizios, Osteoblast adhesion on nanophase ceramics, Biomaterials 20 (13) (1999) 1221–1227. Copyright (1999), with permission from Elsevier.

Modern techniques of surface geometry modification

119

%BIC 100 90 80 70 60 50 40 30 20 10 0 00

4 weeks * 76 53 26

28

m

n 00

3

1)

3)

nm

10

(1:

m

n 00

(1:

34

m

0n

15

AE

G

3

Fig. 2.2.4 Titanium samples obtained by lithography using RIE with a net step of 1000 nm (A), 300 nm—relation ridge/groove ¼ 75/225 nm (B), 300 nm (C), and 150 nm (D) and corresponding BIC parameters after 4 weeks of implantation [21]. Reprinted from L. Prodanov, E. Lamers, M. Domanski, R. Luttge, J.A. Jansen, X.F. Walboomers, The effect of nanometric surface texture on bone contact to titanium implants in rabbit tibia, Biomaterials 34 (12) (2013) 2920–2927. Copyright (2013), with permission from Elsevier.

(GAE
sandblasting + etching) is one of the most efficient implants used at the moment [22,23]. In Ref. [24], the efficiency of the smaller scale relief is shown. A mesoporous TiO2 xerogel layer with a pore diameter of 6 nm was coated on the implant surface by the sol-gel template method. Implantation of these samples leads to an increase of the removal moment after 3 weeks of the engraftment (Fig. 2.2.5). To explain the effect, a test was conducted with the use of simulated body fluid (SBF). This test revealed growth of hydroxyapatite in the pores of the xerogel. Therefore, the nanorelief in addition to mimicry of collagen bundles may affect the mineralization and adhesion area of the implant with biological hydroxyapatite. It is reasonable to propose that the combination of nano- and micron-sized relief in the same implant can significantly improve its biomedical properties. There are some works in this direction [25–29]. The work [30] is exemplary for evaluation of this approach efficiency: a metallic titanium surface was acid-etched in order to create micron-sized relief. This leads to the appearance of micropits. Subsequently, nanonodules were grown by sputtering on the surface. Next, cytological studies were conducted on samples with nodules of different diameters with respect to the acid-etched titanium. Adhesion of cells should be identical due to micron-sized relief of both samples; however, there is evident increase caused by synergy (Fig. 2.2.6). As for the other cytological parameters, a significant difference was also detected (Fig. 2.2.7). On samples with two-level relief hierarchy, the concentration of alkaline phosphatase (that participates in the formation of orthophosphate anions-mineral component of solid tissue) was higher; the content of calcium throughout the period

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20

3 Weeks MP NP

Removal torque (N cm)

18 16 14 12 10 8 6 4 2 0 1

2

3

4

5 6 Rabbit

7

8

9

Mean

Fig. 2.2.5 The results of the RTT test of the implants with or without mesoporous coating [24]. Reprinted from J. Karlsson, R. Jimbo, H.M. Fathali, H.O. Schwartz-Filho, M. Hayashi, M. Halvarsson, M. Andersson, in vivo biomechanical stability of osseointegrating mesoporous TiO2 implants, Acta Biomater. 8 (12) (2012) 4438–4446. Copyright (2012), with permission from Elsevier.

Micropits Micropits + 100-nm nodules Micropits + 300-nm nodules Micropits + 500-nm nodules Cell attachment (WST-1)

Fig. 2.2.6 The dependence of cell adhesion on the type of two-level hierarchy organization of the titanium surface [30]. Reprinted from K. Kubo, N. Tsukimura, F. Iwasa, T. Ueno, L. Saruwatari, H. Aita, T. Ogawa, Cellular behavior on TiO2 nanonodular structures in a micro-tonanoscale hierarchy model, Biomaterials 30 (29) (2009) 5319–5329. Copyright (2009), with permission from Elsevier.

0.1

**

**

**

0.08

* **

0.06

**

0.04 0.02 0

6h

24 h

of incubation was also higher. In addition, the high expression levels of genes responsible for the osteocalcin and collagen production were found. Therefore, we emphasize that the most promising technique of geometric surface modification is the creation of a two-level relief hierarchy. This not only combines the effects of micron and nanoscale topography but also demonstrates a synergistic effect of these two organization levels.

0

Day 7

0.3

0.2

0.1

0

Total calcium deposition (mg/dL)

20

* * Gene expression leve (arbitrary unit)

40

**

Relative ALP activit/ cell

ALP positive area (%)

** 60

*

Day 7

1.2

2.0 1.5

**

** **

* * *

0.9 ** ** **

1.0

0.6

0.5

0.3

0

Day 7

Day 14

Collagen I

0

Day 7

Day 14

6

** **

5 4 3 2 1 0

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**

7 **

80

Day 21

Osteocalcin

Fig. 2.2.7 The dependence of the biochemical parameters of the cells determining the osteosynthesis from the type of two-level hierarchy organization of the titanium surface [30]. Reprinted from K. Kubo, N. Tsukimura, F. Iwasa, T. Ueno, L. Saruwatari, H. Aita, T. Ogawa, Cellular behavior on TiO2 nanonodular structures in a micro-to-nanoscale hierarchy model, Biomaterials 30 (29) (2009) 5319–5329. Copyright (2009), with permission from Elsevier.

121

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Classical methods of the surface geometry modification of titanium implants

2.2.2.1 Mechanical surface treatment The typical purpose of this modification type is to create the implant surface irregularities suitable for the adhesion of osteoblasts. These irregularities reduce the induction period of bone growth on the implant surface. One of the most commonly used methods is abrasive processing. Abrasives are used either in bound form (cemented or sintered abrasive belts, circles and bars) or in free powdered form. Powders for treatment can be administered in the form of suspensions in fibrous materials such as felt. Alternatively, a stream of medium with abrasive is created. For liquid medium, the method is called wet abrasive blasting; for a gaseous medium, it’s called sandblasting. The latter technique is often found in the thematic literature as well as in the praxis because sandblasting is used for the surface treatment of titanium dental implants. In Ref. [31], the authors describe the influence of sandblasting on the biomedical characteristics of titanium implants. Originally titanium samples were processed by sand with different grain sizes (Table 2.2.1). Osteoblasts were incubated on these samples. After the incubation completed, analysis of the alkaline phosphatase concentration was made. The appearance of this enzyme indicates the beginning of bone tissue mineralization. In Fig. 2.2.8, curves of inhibition of alkaline phosphatase are given together with phosphatase activity on the samples with different roughness. In the reviewed work, the conclusion is made that micron-sized roughness has a positive effect. However, the authors base this only on the cytometry methods whereas the microscopic engraftment factors must be confirmed from the point of view of histology. The first argument for confirmation is the growth of the required tissue on the implant’s surface. The second argument (that belongs rather to biomechanics than to histology) is the determination of the removal force moment from the tissue after engraftment. Histological experiments completely confirm the results of the cytology. In Ref. [32], the tissue growth is described on rough and flat surfaces: rough surfaces usually form strong contacts with solid tissues whereas the flat ones could be encapsulated by connecting tissue. Nevertheless, sometimes tissue study is insufficient for a final decision about the efficiency of the surface modification. This is because bone

Table 2.2.1

Abrasive grain size and roughness of titanium samples

No.

Grain size (μm)

Roughness Ra (μm)

1 2 3 4 5

polished 50 75 125 250

0.608 1.012 1.443 2.139 3.217

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123

40 a,* ALP activity (sigma U/mg)

control

a,*

PA(1µM)

30

DPG(10µM) a,* * 20

*

a,* *

a * 10 a

0

0

a

1

a

a

a

2

3

4

Fig. 2.2.8 The dependence of the alkaline phosphatase activity from the roughness of the sand-blasted titanium samples [31]. Reprinted from K. Myung-Joo, C. Myung-Un, K. Chang-Whe, Activation of phospholipase D1 by surface roughness of titanium in MG63 osteoblast-like cell, Biomaterials 27 (2006) 5502–5511. Copyright (2006), with permission from Elsevier.

implantation (especially dental ones) must provide mechanical implants’ holding while loading. Therefore, in addition to the classic study of tissue morphology, additional biomechanical tests are needed that can detect weak contact of the tissue with the implant. Such methods exist and in general they confirm the relationship between roughness and osseointegration. In Ref. [33], flat and sandblasted implants were engrafted into the rabbits’ bone. After 3 months, the removal force moment was measured. For the whole second series this parameter was found to be higher. Therefore, one can conclude that roughness influences positively on the engraftment. However, this is not an unequivocal parameter because several side parameters also occur [34]. In Ref. [35], it is stated that the surface topology description is not unambiguous when using the classic integral value Ra only. This is because it describes only the total surface curvature. At the same time, a full topology description must include also average distance between peaks, average peak heights, etc. Moreover, abrasives often form anisotropic structures (Fig. 2.2.9) [36], which should be additionally described by vector parameters. At the moment, questions about the ideal topology for the healing of the surface generated by machining remain open but the overall trends are generally obvious.

2.2.2.2 Etching Etching is the technology of removal of the surface layer using chemical reagents. There are three main approaches: chemical, electrochemical, and plasma etching. Chemical etching is the oldest known technique. Part of the material is removed

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(A)

(B)

Fig. 2.2.9 Electron microphotographs of the titanium surface: polished (A) and abrasive-processed (B) [36]. Reprinted from J. Lincks, B.D. Boyan, C.R. Blanchard, C.H. Lohmann, Y. Liu, D.L. Cochran, D.D. Dean, Z. Schwartz, Response of MG63 osteoblast-like cells to titanium and titanium alloy is dependent on surface roughness and composition, Biomaterials 19 (1998) 2219–2232. Copyright (1998), with permission from Elsevier.

due to the reaction with etchant—the compound that transforms the sample material into a soluble or gaseous form. Electrochemical etching is the etching of the conducting surface due to the oxidation of the material during redox electrochemical reaction. The sample is used as the dissolvable anode that is placed into an electrolytic solution or melt; cathode is also placed there. After that, current is applied to this cell. As a result, anodic material dissolves in the course of the electrochemical process: М е
е
¼ Me + aq If the reaction proceeds in the diffusion region, roughness is fitted—this is the electrochemical polishing process. Etching in the kinetic region leads to leaching of the porous structure. The third kind of etching is plasma etching: removal of material from the surface by plasma ion bombardment. The most effective approach is reactive ionic etching; this technique uses the combination of etchant, ionic spraying, and ionic activation. It leads to the destruction of the surface layer and to the desorption of volatile reaction products. Let’s consider the application of the reviewed etching techniques for the modification of implant surfaces. Several works were dedicated to the study of acid etching of the titanium surface. The most systematic study was performed in the work [37]. Twenty identical implants (4  3, 25 mm) were prepared. Ten of them were subjected to mechanical treatment; their surface represents the flat with anisotropic micronsized abrasive notches (Fig. 2.2.10), whereas the chemical etching leads to the developed topology of the surface (Fig. 2.2.10B).

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(A)

125

(B)

Fig. 2.2.10 Electronic microphotographs of Ti implants surfaces (А) polished; (B) acid-etched [37]. Republished with permission of John Wiley & Sons, Inc., P.R. Klokkevold, R.D. Nishimura, M. Adachi, A. Caputo, Osseointegration enhanced by chemical etching of the titanium surface: a torque removal study in the rabbit, Clin. Oral Impl. Res. 8 (1997) 442–447. Copyright (1997); permission conveyed through Copyright Clearance Center, Inc.

Acid etching has been performed with aqueous solutions of H2SO4 or HСl. Further, rabbits of the weight of 4.5–5.5 kg were prepared. The flat surface on the lateral distal aspect of the femur was selected for implant placement. Next, holes were prepared using a conical drill. The implants were screwed into the holes, then tissues were restored. After that, the rabbits were kept in vivarium for 2 months with feeding ad libitum. Healing was followed by chemical euthanasia. Subsequently, every implant was cleared surgically and clinically analyzed. Only two implants from 10 mechanically treated ones did not exhibit expected engraftment—they had removal force moments below instrument sensitivity. The remaining 18 implants passed the test successfully. These results are shown in Fig. 2.2.11. Thus, the authors came to the conclusion that etching leads to a high response of bone tissue and, therefore, this is a perspective for the development of implants with improved engraftment. Besides etching with a mixture of sulfuric and hydrochloric acids, other etching agents are also known, showing positive results: Aqueous Aqueous Aqueous Aqueous

solution HF, HCl, H2SO4—[38]. solution HF—[39]. solution HCl, H2O2—[40]. solution HF, HNO3—[41].

The promising application of this approach is confirmed not only by in vitro and in vivo experiments on animals but also by the implantation statistic for real patients [42]. Another important aspect of chemical etching is the appearance of ordered texture on the titanium surface while applying some etching agents. Some microphotographs of the samples etched by the various techniques are given in [45] (Fig. 2.2.12). As one can see, all the etching agents make small holes of various diameters. In the case of the system ammonia-hydrogen peroxide, the net of connected cracks is

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Torque (N cm) 40 35 30 25 20 15 10 5 0 1

Acid etched 2

3

4

5

6

7

Machined 8

9

10

Fig. 2.2.11 Force moments of the removal of implants from the bone [37]. Republished with permission of John Wiley & Sons, Inc., P.R. Klokkevold, R.D. Nishimura, M. Adachi, A. Caputo, Osseointegration enhanced by chemical etching of the titanium surface: a torque removal study in the rabbit, Clin. Oral Impl. Res. 8 (1997) 442–447. Copyright (1997); permission conveyed through Copyright Clearance Center, Inc.

observed. In addition to the effect of the etchant in [43,44] also studied the effect of grain structure of titanium on the surface texture after etching. This is a very interesting direction, but, at the moment, the effect of the texture is poorly understood. Summing up this section, it should be noted that surface etching is the simple method from the practical point of view. Applying the variety of implementation options and a wide array of etchants, it is possible to create a wide range of surface textures, including those suitable for improving the titanium bioactivity.

2.2.2.3 Anodization Electrochemical etching can take place without complete dissolution of anode material in the etchant. In this case, an oxide layer grows on the surface. This is anodization that is widely used for the modification of titanium implant surfaces [45–47]. In Ref. [46], microphotographs of the Ti alloy surface are given after anodization in aqueous sulfuric acid (Fig. 2.2.13). So, one can observe the presence of a nanometer-sized texture on the sample surface. At lower magnification, a weakly developed micron-sized texture can be revealed (Fig. 2.2.14). Therefore, this technique is also prospective for implant development. This method looks even more attractive when considering the works describing the formation of the structures, anisotropic along the normal to the surface. Under certain conditions localization of the charge occurs during the anodization on the areas with a high etching

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Fig. 2.2.12 Electronic microphotographs of the samples, treated by different etching agents [45]. Reprinted from F. Vetrone, F. Variola, P. Tambasco de Oliveira, S.F. Zalzal, J.-H. Yi, J. Sam, K.F. Bombonato-Prado, A. Sarkissian, D.F. Perepichka, J.D. Wuest, F. Rosei, A. Nanci, Nanoscale oxidative patterning of metallic surfaces to modulate cell activity and fate, Nano Lett. 9 (2) (2009) 659–665. Copyright (2009) American Chemical Society.

rate. As a result, these areas repelled and formed a distorted hexagonal pore network [48] (Fig. 2.2.15). Such a structure has good cellular response. So, in Ref. [49] the same nanotubes were obtained on the originally etched surface of titanium. These structures had no significant effect on osteoblast proliferation; however, a huge increase of alkaline

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Fig. 2.2.13 Electronic microphotograph of an anodized titanium alloy surface [46]. Reprinted from N. Masahashi, Y. Mizukoshi, S. Semboshi, K. Ohmura, S. Hanada, Photo-induced properties of anodic oxide films on Ti-6Al-4V, Thin Solid Films 520 (2012) 4956–4964. Copyright (2012), with permission from Elsevier.

Fig. 2.2.14 Electronic microphotograph of anodized titanium alloy surface [46] at low magnification. Reprinted from N. Masahashi, Y. Mizukoshi, S. Semboshi, K. Ohmura, S. Hanada, Photoinduced properties of anodic oxide films on Ti-6Al-4V, Thin Solid Films 520 (2012) 4956–4964. Copyright (2012), with permission from Elsevier.

phosphatase activity was discovered, indicating a significant promotion of the differentiation process. Subsequent studies showed that this effect on the surfaces after the classical anodizing is less pronounced [50]. Thus, it is concluded that the anodizing is suitable for creating bioactive coatings. The use of pile-like surfaces is especially efficient.

2.2.2.4 Coating of TiO2-based materials The previous sections have been devoted to the surface texturing techniques by means of removing part of the implant material. In this section we will consider methods of applying coatings to titanium, which allows us to obtain a surface with a given topology without affecting the bulk implant. Coating methods are divided into physical and chemical ones. In the first case, these are the processes such as mechanical drawing, condensation from gas, and deposition from the liquid. As one can see, the substance does not undergo composition changes in the course of coating. In the second group of techniques, the coating substance has a different composition than the one used to obtain the initial reagents. There are several such methods: application of lacquer

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(A)

129

(B)

Fig. 2.2.15 Electron microphotograph of anodized porous (tubular) titanium surface (A). Electron microphotograph of a chip of the surface (B) [48]. Reprinted from M. Goudarzi, F. Batmanghelich, A. Afshar, A. Dolati, G. Mortazavi, Development of electrophoretically deposited hydroxyapatite coatings on anodized nanotubular TiO2 structures: Corrosion and sintering temperature, Appl. Surf. Sci. 301 (2014) 250–257. Copyright (2014), with permission from Elsevier.

and paint coatings with polymerization in a thin layer, chemical deposition from the gas phase, sol-gel technology, etc. In these processes, physical methods are also required—at least for transport onto a substrate surface. Therefore, chemical methods are more complicated while also requiring both development of the necessary reaction and the way of localization of the reaction on the surface. Every approach has its own advantages and drawbacks. Let’s consider them in more detail.

2.2.2.4.1 Physical methods of applying titanium and titanium oxide-based texturing coating for increasing implant bioactivity 2.2.2.4.1.1

Gas thermal spraying

There are a number of works devoted to the production of titanium coatings by thermal spraying. At first glance, the deposition of a titanium layer on the titanium surface does not make sense. In fact, though, this technique is very promising for faster osseointegration as precipitation may take place either in a continuous film growth mode or in the mode of formation of defined irregularities that may have a cellular response. This method includes heating, dispersing, and transfer of the deposited particles of the sputtered material by gas flow that leads to the formation of the deposited layer on the substrate. The practical implementation is quite simple: a burner is used that flame transfers metal particles. Despite its simplicity, this method is very effective because deposited titanium particles weld to the implant. This results in a high adhesion and thus the particles partially retain their original shape that leads to the textured surface (Fig. 2.2.16). In the work, the samples of sputtered titanium oxide were obtained on titanium from powders with different grain sizes. During the cytological measurement, it

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Fig. 2.2.16 Surface texture obtained by gas thermal spraying of titanium oxide [53]. Reprinted from W. Xue, X. Liu, X.B. Zheng, C. Ding, in vivo evaluation of plasma-sprayed titanium coating after alkali modification, Biomaterials 26 (2005) 3029–3037. Copyright (2005), with permission from Elsevier. NONE

SEI

10.0 kV ×10,000

1 µm WD 7.7 mm

was found that cell proliferation occurs faster on the modified surface. The strong dependence of the effect was also noted with the variation of grain size of powder particles and deposition temperature.

2.2.2.4.1.2

Physical vapor deposition

There are several approaches to the implementation of this technique: the closest to the above-mentioned thermal spraying method is an atmospheric plasma spraying. Here, coating material is introduced into the plasma jet created by an electric arc where material itself is undergoing ionization. Coating in an inert atmosphere is preferable because it excludes the reaction of the sprayed material with air. Usually this variation is called plasma spraying in a controlled atmosphere or an inert plasma spraying. Sputtering is applied when one needs to produce the coating under mild conditions. This technique is based on bombardment of the sputtering target with ions. As the result, ions knocked out from the material and subsequently condensed on the substrate. For obtaining such conditions, the plasma of the inert gas can be applied that is usually produced by a magnetron or high electric potential (cathode sputtering). Sputtering can be used to create a composite implant due to the small influence on the substrate, as it allows creating a surface similar to a bulk titanium, its alloys, and even materials not related to the titanium. Also, for sputtering targets in a vacuum, a resistance furnace can be used (thermal spraying), laser (laser ablation), and even some less-common techniques. Sputtering techniques can work both in a continuous film growth mode and in the granular film growth mode (Fig. 2.2.17) [51], and even for anisotropic objects (Fig. 2.2.18). The second type of surface is very interesting and partially similar to the texture of the tubular anodized surface [49]. In Ref. [52], data are shown of the cell sorption and acceleration of their differentiation on this surface as well as occurrence of abnormally small pseudopodia of osteoblasts that may increase adhesion of bone tissue with the surface. In addition, the authors emphasized the significant volume of cavities suitable

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Fig. 2.2.17 The structure of titanium coating produced by vacuum plasma spraying [51]. In the first case we are dealing with the surface, similar to the classic anodic treatment. Reprinted from D. Weingart, S. Steinemann, W. Schilli, J.R. Strub, U. Hellerich, J. Assenmacher, J. Simpson, Titanium deposition in regional lymph nodes after insertion of titanium screw implants in maxillofacial region, Int. J. Oral Maxillofac. Surg. 23 (1994) 450–452. Copyright (1994), with permission from Elsevier.

Vertical columns

Fig. 2.2.18 Electron micrograph of pile-like titanium coating produced by plasma deposition at grazing angles [52]. Reprinted from Y. Motemani, C. Greulich, C. Khare, M. Lopian, P.J.S. Buenconsejo, T.A. Schildhauer, A. Ludwig, M. K€oller, Adherence of human mesenchymal stem cells on Ti and TiO2 nano-columnar surfaces fabricated by glancing anglesputter deposition, Appl. Surf. Sci. 292 (2014) 626–631. Copyright (2014), with permission from Elsevier.

for filling a variety of drugs. It should be noted that the surfaces having pores [45,46,38,51] and tubes [49] are also suitable for filling but this aspect will not be the object of our review. In addition to texturing, the titanium-based coating can be a barrier to the toxic ion emissions from the bulk implant. Thus, in Ref. [53] an implant made of Ti-6Al-4V alloy has better mechanical properties than pure titanium but inferior biocompatibility due to the presence of vanadium. It was coated with a titanium layer with a thickness of 100 μm by plasma spraying in a controlled atmosphere. As a result, this surface shows good qualities during cytological experiments.

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In addition to these physical deposition methods, there are many others but they are not suitable either for the deposition of titanium coating or for depositing on a titanium substrate. However, in the medical literature associated with titanium implants, the most common method is spraying.

2.2.2.5 Chemical methods of applying texturing coating based on titanium oxide to increase the bioactivity of implants 2.2.2.5.1 Chemical vapor deposition The physical vapor deposition that we have just discussed has a chemical analogue: chemical vapor deposition or CVD. This chemical technique is using for producing both thin film and massive objects as well as for obtaining very pure substances. In the case of the thin-film application of this technique, the substrate is placed in a pair of one or more substances that during mutual reaction or decomposition is deposited as a product layer on the surface. The reaction products are usually removed from the reactor and the surface of the treated substrate by carrier gas or vacuum. However, despite the proximity of the physical and chemical methods of deposition, the search of works on obtaining coatings by CVD has revealed a very weak development in the area in general: a small number of publications and lack of a systematic approach to the analysis of biological parameters. So in the article [54] for TiO2 coating on dental implants obtained by CVD, the antibacterial properties were analyzed while neither analysis of the survival cell nor cytological studies had been performed. Thus, we must conclude that despite the prospects of applying the CVD method to create nanotexture on the surface of titanium implants, this area remains up to now almost unexplored. Most of the examples of this method of application are associated not with the geometry of the oxide but with the introduction of additional compounds in a surface titanium layer.

2.2.2.5.2 The sol-gel technique Another method of creating a textured coating on the implant surface is sol-gel technology. As with the previous method, it refers to the chemical: it consists of the formation of solid material or coating by condensation during the sol-gel process. Relating to the considered technology, the gels prepared from sols are a grid of condensed nanoparticles filled with mother liquor also containing products of the condensation reaction. Removal of the solution (dispersion medium) leads to the so-called xerogels—solid bodies composed of cross-linked sol particles. This coating typically has a high porosity due to the presence of the unfilled internal space in the frame after the removal of the dispersion medium. A distinctive feature of the method is the precipitation of gel in the so-called “soft chemistry,” meaning reactions occur at low

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133

temperatures and, as a rule, do not require any other special conditions such as a vacuum, pressurized activation initiators, radiation, etc. This makes the technique very attractive from the point of view of technological applications. There are a number of articles describing the production of monolithic [55], lowporosity titanium oxide films by the sol-gel method. However, their application to create biocompatible coatings for titanium implants does not make sense because they are not much different from the natural titanium oxide. Works that are more interesting are based on the creation of porous films. In Ref. [56] there is described a very simple but original method of producing a titanium coating by the sol-gel method revealing a porous film of a hydrated titanium oxide. The coating was deposited from sol aged at an elevated temperature by hydrolysis of titanium tetraisopropoxide in an acidic medium. For the application, a very common method is used—dip coating. This method consists of extracting the substrate at a certain speed from the solution, resulting in a film deposited on its surface. The obtained result was a uniform layer of porous oxide. Cytological studies showed that the coatings have good cell response; these coatings also have pores suitable for introducing different osseointegration promoters. In the case of the deposition of gels, simple condensation of the sol particles is possible to obtain a porous structure with a wide pore size distribution and irregular geometries. A templating method is used to create ordered structures. In the case of the sol-gel technique, it is based on the introduction of various macromolecules, aggregates, or nano objects into the gel during the condensation; its subsequent removal results in xerogels with the pores repeating the object form. In Ref. [24], the authors synthesized mesoporous titania film. Synthesis was carried out by hydrolysis of titanium tetraisopropoxide in an acidic medium in the presence of a nucleating template—micelle of surfactant Pluronic P123. At concentrations above second critical micelle-forming concentration, this material forms cylindrical micelles that can be built in spontaneously hexagonalpacked anisotropic liquid crystal. The authors used this feature: the gel deposited in the presence of Pluronic P123 was a composite of the oxide matrix filled with an organic component and the material itself repeats the structure of the liquid crystal. After removing the surfactant, residues, and mother liquor from the composite, xerogels with cylindrical hexagonal-packed pores were formed. For film deposition, the authors chose a frequently used method—spin coating based on applying the solution on a rotating substrate resulting in a spreading of the solution from its center under the centrifugal force. Fig. 2.2.19 shows the electron microphotographs of the resulting film. On the first picture (Fig. 1.19A), we can see an ordered porous structure of resulting film. On the second (Fig. 1.19B), we see a packless due to the large volume of pore structure of the layer cut through to the substrate and the nanometer pore throat on the surface. The resulting films were tested both in vitro and in vivo. As a result, it was found that the implant extraction torque is higher than for the nonporous surface, especially at a low engraftment timing that is particularly important because during this time the contact with the implant (BIC) is particularly low.

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50 nm

(A)

100 nm

(B)

Fig. 2.2.19 Electron microphotographs of the mesoporous films of titanium oxide: transmission (A) and scanning electron microscopy (B) [24]. Reprinted from J. Karlsson, R. Jimbo, H.M. Fathali, H.O. Schwartz-Filho, M. Hayashi, M. Halvarsson, M. Andersson, in vivo biomechanical stability of osseointegrating mesoporous TiO2 implants, Acta Biomater. 8 (12) (2012) 4438–4446. Copyright (2012), with permission from Elsevier.

2.2.3

Prospective methods of geometry implant surface changes to create a two-level hierarchy of topography

The above-considered classical modification techniques can be used to create a twolevel hierarchy of the implant surface topography. This requires a combination of methods of the micron and nanometer topography creation. In the first part of the review, we have already mentioned the combination of acid etching and cathode sputtering methods [30]: etching leads to the micron-sized pitting, whereas sputtering produces nanometer-sized titanium nodules (Fig. 2.2.20). Fig. 2.2.20 Electron microphotograph of the structure of the titanium surface after etching and deposition nodules by cathodic sputtering [30]. Reprinted from K. Kubo, N. Tsukimura, F. Iwasa, T. Ueno, L. Saruwatari, H. Aita, T. Ogawa, Cellular behavior on TiO2 nanonodular structures in a micro-to-nanoscale hierarchy model, Biomaterials 30(29) (2009) 5319–5329. Copyright (2009), with permission from Elsevier.

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Such a surface structure significantly increases the bioavailability of implants. In Ref. [25] a combination of dual acid etching with anodic treatment was used. This led to the growth of nanotubes normally oriented to the surface (Fig. 2.2.21). As expected, such organization of the surface layer will provide a significant cytological response. However, the author has carried out even more interesting research in comparison with the majority of related works: in addition to the classical pure titanium surfaces, samples after etching and anodic treatment used as standard. As a result, it was clearly shown that the effect has nanometer and micron relief but their combination in the same sample has a strongly marked synergy (Fig. 2.2.22). Next, we will skip the careful review of bioactivity studies of different types of surfaces with a two-level structure of the relief, as in all presented papers a positive effect of the modification is observed. Let us consider the techniques of the relief creation. In Ref. [26], using the combination of acidic and alkaline etching, so-called nest-like structures were obtained with a very large surface area (Fig. 2.2.23). Another combination of surface treatment methods was used in Ref. [29]: micronsized roughness achieved by sandblasting whereas nanometer-sized roughness—by alkaline treatment. This led to the appearance of the titanium oxide nanoplates (Fig. 2.2.24). Another technologically advanced and promising approach of a two-level surface relief creation is presented in Ref. [57]. To create nanorelief, the above-mentioned anodic oxidation was applied. This leads to the formation of titanium-oxide nanotubes and, in contrast to previous works, micron-sized relief was achieved without erosion of the metal substrate. The titanium was processed by the vacuum plasma spray instrument. Titanium powder was applied as the starting material, resulting in the formation of the developed micron relief (Fig. 2.2.25).

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Fig. 2.2.21 Low-resolution (A) and high-resolution electron microphotographs (B) of the titanium surface subjected to acid etching and anodic treatment [25]. Republished with permission of John Wiley & Sons, Inc., X. Chen, K. Cai, M. Lai, L. Zhao, L. Tang, Mesenchymal stem cells differentiation on hierarchically micro/nano-structured titanium substrates, Adv. Eng. Mater. 14 (5) B216–B223. Copyright (2012); permission conveyed through Copyright Clearance Center, Inc.

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Fig. 2.2.22 Integral values of protein adsorption (A), the number of cells (B), alkaline phosphatase activity (C), and normalized mineralization (D) for pure titanium, nanotube-coated titanium, nanotube-coated etched titanium, and etched titanium [25]. Republished with permission of John Wiley & Sons, Inc., X. Chen, K. Cai, M. Lai, L. Zhao, L. Tang, Mesenchymal stem cells differentiation on hierarchically micro/nano-structured titanium substrates, Adv. Eng. Mater. 14 (5) B216–B223. Copyright (2012); permission conveyed through Copyright Clearance Center, Inc.

20 µm

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Fig. 2.2.23 Low-resolution (A) and high-resolution electron microphotographs (B) of the titanium surface subjected to acid etching and alkaline treatment [26]. Reprinted from P. Jiang, J. Liang, C. Lin, Construction of micro-nano network structure on titanium surface for improving bioactivity, Appl. Surf. Sci. 280 (2013) 373–380. Copyright (2013), with permission from Elsevier.

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Fig. 2.2.24 Electron microphotograph of the titanium surface after sandblasting, alkaline treatment, and heat treatment [29]. Reprinted from T. Ueno, N. Tsukimura, M. Yamada, T. Ogawa, Enhanced bone-integration capability of alkali-and heat-treated nanopolymorphic titanium in micro-to-nanoscale hierarchy, Biomaterials 32 (30) (2011) 7297–7308. Copyright (2011), with permission from Elsevier.

400 µm (A)

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Fig. 2.2.25 Low-resolution (A) and high-resolution electron microphotographs (B) of the titanium surface subjected to plasma spraying followed by titanium anodic treatment [57]. Reprinted from Y. Xie, H. Ao, S. Xin, X. Zheng, C. Ding, Enhanced cellular responses to titanium coating with hierarchical hybrid structure, Mater. Sci. Eng. C 38 (2014) 272–277. Copyright (2014), with permission from Elsevier.

Another method that does not affect the implant material is texturing the film during the heat treatment. In Ref. [58], TiO2 gel film deposited on the titanium surface was subjected to shock drying, resulting in shrinkage due to the loss of the dispersion medium. To compensate tensions, the micron cracks were formed on the film. Nanorelief was formed during the synthesis of the xerogel from the gel (Fig. 2.2.26). We have considered a wide range of coating methods for creating a two-level relief organization. They all are united by a combination of methods that provide micron and nanometer relief. This is a combination of sandblasting and anodic treatment, acid and alkali etching, sol-gel synthesis and shock drying, acid etching and cathode sputtering, etc. However, in the literature an approach was found that created two levels of relief at a time just by one method. In Ref. [27] the femtosecond laser treatment resulted in the formation of micron hills with nanometer roughness of the outer surface (Fig. 2.2.27).

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14.65 nm 18.55 nm 8.789 nm

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100 nm

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Fig. 2.2.26 Low-resolution (A) and high-resolution (B) electron microphotographs of the titanium surface coated with TiO2 xerogel obtained by shock drying [58]. Reprinted from E.G. Zemtsova, A.Y. Arbenin, R.Z. Valiev, E.V. Orekhov, V.G. Semenov, V.M. Smirnov, Two-level micro-to-nanoscale hierarchical TiO2 nanolayers on titanium surface, Materials 9 (12) (2016) 1010 with permission from Zemtsova, E.G.

Fig. 2.2.27 Low-resolution (A) and high-resolution electron micrographs (B) of the titanium surface treated with the femtosecond laser [27]. Republished with permission of John Wiley & Sons, Inc., J.R. Bush, B.K. Nayak, L.S. Nair, M.C. Gupta, C.T. Laurencin, Improved bio-implant using ultrafast laser induced self-assembled nanotexture in titanium, J. Biomed. Mater. Res. B: Appl. Biomater. 97 (2) (2011) 299–305. Copyright (2011); permission conveyed through Copyright Clearance Center, Inc.

30.0 µm

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Thus, now there are many methods of titanium surface modification that allow us to create a relief with a two-level hierarchy structure. The vast majority of publications discuss the techniques based on standard methods of synthesis: anodic treatment [59], termooxidizing [60], acid etching [61], laser processing [62], sandblasting [63], etc. Despite the difficulty of creating a two-level hierarchy of structures that combine the lower level, which is in the range of 1–1000 nm, and the upper level in the range of 1–100 μm, this problem is, in terms of technology, solved. Therefore, we must conclude that the approach to improve bone implant bioactivity is promising not only due to increased biomedical characteristics but also because of the processability of the used modification methods.

2.2.4

The practical application of the coating with a two-level hierarchy of the surface relief in implantology

In terms of the subject, just in 2016 we found a dozen articles describing new methods of synthesis and properties of bioactive coatings with a two-level hierarchy [64–72,58] that explicitly refer to the active development of the area. This can be explained by the fact that the existing implants are far from ideal. This also applies to the mechanical properties of metallic bases: in the direction of active work on the creation of new alloys [73,74], ultrafine metal [75,76] and implant designs, work on the creation and application of methods does not stop publication for several decades [77–80]. This is a very important direction but it should be noted that in various fields of implantation, the main problem is the interface of the implant/biological tissues: the rejection [81], infection of preimplanting region [82], structure change of preimplanting field during continuous exploitation: periimplantit [83], hydroxyapatite bioresorption [84], the leaching of toxic metals from medical titanium alloys [85], and so on. Solving these problems often requires the surface modification of implants and one of the most promising ways is the creation of a two-level hierarchy of relief because such coatings, as stated above, accelerate engraftment. By accelerating this process, we can reduce the possibility of implant rejection and infection of the periimplantit preimplanting area; the coating can block the release of toxic ions. One more exciting possibility of the developed surface is the transfer of additional useful components: antibiotics, osseointegration promoters such as calcium phosphates, bisphosphonates, and so on. Thus, we must conclude that this is a promising area of implantology due to an opportunity to improve the properties of biomedical implants and manufacturability of applying approaches.

2.2.5

Conclusion

Summarizing our review, we should note that periimplantit long-term experience of titanium bone implant applications revealed a significant influence of geometrical surface modifications on biomedical properties. This is the influence of nano- and

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microrelief on such processes as the differentiation of osteoblasts accelerating, deposition of biological hydroxyapatite in nanometer roughness and blocking the osteoblasts drift in micron-sized roughness. But work over recent decades demonstrates that the optimal is the combination of two types of relief in one surface layer of the implant. This allows the combination of effects peculiar to each type and gives a synergistic effect. A two-level relief hierarchy allows the reduction of the healing period and the improvement of the adhesion of bone tissue with the implant. This will significantly improve the quality of implants in terms of patient comfort while also reducing the statistics of tissue injury and implant loss. Most of the work in this area describes the application of well-known methods for creating a two-level relief hierarchy, which indicates the possibility of their practical implementation. Thus, we can conclude that this is one of the most promising areas in bone implantation.

Acknowledgments The authors acknowledge the Russian Ministry for Education and Science for funding through Contract No. 14.B25.31.0017 by 28 June 2013. This work was also supported in part by a grants from the St-Petersburg State University, No. 12.37.433.2015 and 6.38.337.2015.

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[60] G. Wang, J. Li, K. Lv, W. Zhang, X. Ding, G. Yang, et al., Surface thermal oxidation on titanium implants to enhance osteogenic activity and in vivo osseointegration, Sci. Rep. 6 (2016) 1–13. [61] S.F. Lamolle, M. Monjo, M. Rubert, H.J. Haugen, S.P. Lyngstadaas, J.E. Ellingsen, The effect of hydrofluoric acid treatment of titanium surface on nanostructural and chemical changes and the growth of MC3T3-E1 cells, Biomaterials 30 (5) (2009) 736–742. [62] G. Wang, J. Li, K. Lv, W. Zhang, X. Ding, G. Yang, et al., Surface thermal oxidation on titanium implants to enhance osteogenic activity and in vivo osseointegration, Sci. Rep. 6 (2016). [63] Y. Li, S. Zou, D. Wang, G. Feng, C. Bao, J. Hu, The effect of hydrofluoric acid treatment on titanium implant osseointegration in ovariectomized rats, Biomaterials 31 (12) (2010) 3266–3273. [64] I.H. Kim, J.S. Son, S.H. Choi, K.H. Kim, T.Y. Kwon, Nano-and micro-scale oxidative patterning of titanium implant surfaces for improved surface wettability, J. Nanosci. Nanotechnol. 16 (2) (2016) 1883–1886. [65] J.Y. Xu, X.S. Chen, C.Y. Zhang, Y. Liu, J. Wang, F.L. Deng, Improved bioactivity of selective laser melting titanium: Surface modification with micro-/nano-textured hierarchical topography and bone regeneration performance evaluation, Mater. Sci. Eng. C 68 (2016) 229–240. [66] C. Yin, Y. Zhang, Q. Cai, B. Li, H. Yang, H. Wang, et al., Effects of the micro-nano surface topography of titanium alloy on the biological responses of osteoblast, J. Biomed. Mater. Res. A 105 (3) (2016) 757–769. [67] S.L. Hyzy, A. Cheng, D.J. Cohen, G. Yatzkaier, A.J. Whitehead, R.M. Clohessy, et al., Novel hydrophilic nanostructured microtexture on direct metal laser sintered Ti-6Al4V surfaces enhances osteoblast response in vitro and osseointegration in a rabbit model, J. Biomed. Mater. Res. A 104 (8) (2016) 2086–2098. [68] P. Yu, X. Zhu, X. Wang, S. Wang, W. Li, G. Tan, et al., Periodic nanoneedle and buffer zones constructed on a titanium surface promote osteogenic differentiation and bone calcification in vivo, Adv. Healthc. Mater. 5 (3) (2016) 364–372. [69] D. Kuczynska, P. Kwasniak, J. Marczak, J. Bonarski, J. Smolik, H. Garbacz, Laser surface treatment and the resultant hierarchical topography of Ti grade 2 for biomedical application, Appl. Surf. Sci. 390 (2016) 560–569. [70] G. Li, H. Cao, W. Zhang, X. Ding, G. Yang, Y. Qiao, et al., Enhanced osseointegration of hierarchical micro/Nanotopographic titanium fabricated by microarc oxidation and electrochemical treatment, ACS Appl. Mater. Interfaces 8 (6) (2016) 3840–3852. [71] F.A. Shah, M.L. Johansson, O. Omar, H. Simonsson, A. Palmquist, P. Thomsen, Lasermodified surface enhances osseointegration and biomechanical anchorage of commercially pure titanium implants for bone-anchored hearing systems, PLoS One 11 (6) (2016), e0157504. [72] Q. Fu, A. Bellare, Y. Cui, B. Cheng, S. Xu, L. Kong, The effect of hierarchical micro/ nanotextured titanium implants on osseointegration immediately after tooth extraction in beagle dogs, Clin. Impl. Dentist. Rel. Res. (2016). [73] X. Meng, X. Wang, Y. Guo, S. Ma, W. Luo, X. Xiang, et al., Biocompatibility evaluation of a newly developed Ti-Nb-Zr-Ta-Si alloy implant, J. Biomater. Tissue Eng. 6 (11) (2016) 861–869. [74] K. Hagihara, T. Nakano, H. Maki, Y. Umakoshi, M. Niinomi, Isotropic plasticity of β-type Ti-29Nb-13Ta-4.6Zr alloy single crystals for the development of single crystalline β-Ti implants, Sci. Rep. 6 (2016) 1–10.

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[75] I.P. Semenova, G.V. Klevtsov, N.Y.A. Klevtsova, G.S. Dyakonov, A.A. Matchin, R.Z. Valiev, Nanostructured titanium for maxillofacial mini-implants, Adv. Eng. Mater 18 (7) (2016) 1216–1224. [76] B. An, Z. Li, X. Diao, H. Xin, Q. Zhang, X. Jia, et al., In vitro and in vivo studies of ultrafine-grain Ti as dental implant material processed by ECAP, Mater. Sci. Eng. C 67 (2016) 34–41. [77] A. Blum, J.B. Meyer, A. Raymond, M. Louis, O. Bakour, R. Kechidi, et al., CT of hip prosthesis: new techniques and new paradigms, Diagnos. Intervent. Imaging 97 (7) (2016) 725–733. [78] S.S. Porter, Platform switching: a new concept in implant dentistry for controlling postrestorative crestal bone levels, Dent 26 (2006) 9–17. [79] H. Wehrbein, J. Glatzmaier, U. Mundwiller, P. Diedrich, The orthosystem—a new implant system for orthodontic anchorage in the palate, J. Orofac. Orthoped. Fortschr. Kieferorthopad.: Organ/Off. J. Deutsche Gesellsch. Kieferorthopad. 57 (3) (1996) 142–153. [80] J.S. Hirschhorn, A.A. McBeath, M.R. Dustoor, Porous titanium surgical implant materials, J. Biomed. Mater. Res. 5 (6) (1971) 49–67. [81] R.J. Weyant, Characteristics associated with the loss and peri-implant tissue health of endosseous dental implants, Int. J. Oral Maxillofac. Impl 9 (1) (1994) 95–102. [82] H. Tsuchiya, T. Shirai, H. Nishida, H. Murakami, T. Kabata, N. Yamamoto, et al., Innovative antimicrobial coating of titanium implants with iodine, J. Orthop. Sci. 17 (5) (2012) 595–604. € [83] T. Albrektsson, B. Chrcanovic, P.O. Ostman, L. Sennerby, Initial and long-term crestal bone responses to modern dental implants, Periodontology 73 (1) (2017) 41–50. [84] S. Wenisch, J.P. Stahl, U. Horas, C. Heiss, O. Kilian, K. Trinkaus, et al., In vivo mechanisms of hydroxyapatite ceramic degradation by osteoclasts: fine structural microscopy, J. Biomed. Mater. Res. A 67 (3) (2003) 713–718. [85] N.J. Hallab, S. Anderson, M. Caicedo, A. Brasher, K. Mikecz, J.J. Jacobs, Effects of soluble metals on human peri-implant cells, J. Biomed. Mater. Res. A 74 (1) (2005) 124–140.

Nanobioceramic thin films: Surface modifications and cellular responses on titanium implants

2.3

A.H. Choi*, S. Akyol†, A. Bendavid‡, B. Ben-Nissan* *University of Technology Sydney, Ultimo, NSW, Australia, †University of Istanbul, Istanbul, Turkey, ‡Commonwealth Scientific and Industrial Research Organisation, Lindfield, NSW, Australia

2.3.1

Introduction

Widely accepted as a biocompatible material similar to the mineral components of teeth and bone, hydroxyapatite (HAp) has been utilized as a porous coating material on a range of metallic implants for dental and orthopedic applications since the early 1980s [1–5]. HAp, according to most published data, is categorized as calcium phosphate where it belongs. For this reason, the chemical properties will be considered from the standpoint that HAp is calcium phosphate, even though it will have different solubility and reactivities to other phosphates within the physiological environment. Calcium phosphates are classified by particular solubilities within the body and their bonding mechanisms to the surrounding tissues, and hence by their ability to degrade [6,7]. As calcium phosphate (or HAp) comes into contact with bodily fluid, its surface ions can be exchanged with those of the aqueous solution. Alternatively, various ions and molecules such as collagen and proteins can be adsorbed onto their surface to produce biological films and coatings [8,9]. Despite the fact that bulk calcium phosphates have been utilized since the 1920s for surgical applications as macrocoatings [10], microcoatings and nanocoatings of calcium phosphate are comparably new and were used since the early 1980s as porous coating materials [11]. During the last three decades, issues and concerns arose from the relationship between adhesion and biological interactions. An increase in the reliability and longevity of implant systems has driven the exploration of surface modification toward bone-implant adaptability, rapid healing, and early osteointegration. The main focus for a large number of researchers in the surgical and biomedical fields has been concentrated on improving implant surfaces through macro- and microtexturing or by increasing the bioactivity and interconnectivity via biological and chemical means. The main intention of using calcium phosphate as a bioactive coating is to establish a strong and rapid biological attachment to the soft or hard tissue. The term biological Titanium in Medical and Dental Applications. https://doi.org/10.1016/B978-0-12-812456-7.00007-X © 2018 Elsevier Inc. All rights reserved.

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fixation refers to the process where the implant or prosthetic component firmly bonds to the host tissue without the use of any adhesive but instead through bone in-growth with or without the assistance of mechanical fixation. In general, two different approaches are currently being used to surgically insert orthopedic implants. The first approach utilizes bone cements, mostly poly(methyl methacrylate) (PMMA), for strong adhesion while the second approach uses calcium phosphate precoated porous or microtextured implants for mechanical interlocking and chemical bonding. In addition to orthopedic applications, the latter approach is also widely employed in dental and maxillofacial implants. The benefit of using calcium phosphate coatings rests on the supply of calcium and phosphate ions that may facilitate and/or stimulate the growth of new bone tissue on and toward the implant. A good balance, on the other hand, must be attained in the rate of bone growth in order to achieve mechanical integrity under functional loading and the solubility of the coating. This would allow adequate mechanical properties and bonding at the tissue-implant interface in an effort to achieve long-term survival of an implant or prosthesis [8,9,12]. It has been suggested that, for a successful clinical outcome of a biocompatible coated implant, the ideal environment for bone growth is achieved with a decrease in the release of the metal ion as well as the availability of a bioactive surface for biological and chemical bonds and the creation of good mechanical interlocking [5]. Implants coated with calcium phosphate have been demonstrated to display extensive bone apposition in animal models. The development of good interfacial strength between the bone tissue and the implant is the consequence of the biological interactions of released calcium and phosphate ions. Manufactured in an acceptable manner, calcium phosphate-coated implants heal faster and demonstrate improvements in bone attachment. Factors governing the quality and long-term performance of a calcium phosphate-coated implant include the crystallinity of the coating, constituent phases, porosity, surface roughness, and thickness in addition to the biomechanical functional loading as well as the overall design of the implant and/or prosthesis [8,9,12]. Moreover, the surface topography and chemistry of calcium phosphate crystals deposited as thin film or coating on implants demonstrate an increase in bond strength between bone and implant and acceleration in early bone formation. On the other hand, the properties and microstructures of nanostructured materials such as nanoparticles, nanocoatings, and nanocomposites are governed by factors such as their synthesis or manufacturing methods, chemical composition, structure, and thickness of the coating. Nanocoatings are thin films that may contain homogeneous and isotropic compounds with thicknesses below the range of micron-sized coatings (i.e., 1000 patients had benefited from this development. The other is the world’s first 3D artificial vertebral body, approved by the CFDA in May 2016 [4]. More additively manufactured orthopedic implants are in the process of being assessed for approval by the CFDA, which normally takes 3–5 years. Titanium and titanium alloys have many advantages as surgical implant materials, including relatively low density, high strength, good toughness, good biocompatibility, low elastic modulus, excellent corrosion resistance, and low X-ray absorption rate. Hence, they occupy an increasingly important market share of medical metal materials. There are two grades of Ti-6Al-4V. Grade 23 extra low interstitial (ELI) Ti-6Al-4V is the premier titanium alloy used for medical applications. Its composition is listed in Table 3.3.1 together with that of Grade 5 Ti-6Al-4V [5]. According to ISO5832-3 titanium alloy medical standards [6], both Grade 5 and Grade 23 are in line with the requirements for medical applications in terms of their composition. Grade 23 ELI Ti-6Al-4V offers better ductility but lower strength than Grade 5 Ti-6Al-4V, and can be made into a wide variety of product forms for orthopedic applications. With the increasing acceptance of metal AM, it is common practice today to produce Ti-6Al-4V implants (knee, hip, neck, shoulder, spine, heel, skull, etc.) by AM. Compared to other metal AM processes, SEBM is particularly suited to the manufacture of titanium and titanium alloys due to its high vacuum environment. Titanium in Medical and Dental Applications. https://doi.org/10.1016/B978-0-12-812456-7.00011-1 © 2018 Elsevier Inc. All rights reserved.

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Table 3.3.1 Composition of Ti-6Al-4V in Grade 5 and Grade 23 (extra low interstitial) Other impurities total(each)

Grade

Al

V

O

Fe

C

H

N

5

5.5–6.75

3.5–4.5

0.20

0.30

0.08

0.015

0.05

23(ELI)

5.5–6.5

3.5–4.5

0.13

0.25

0.08

0.0125

0.03

ISO 5832–3

5.5–6.75

3.5–4.5

0.20

0.30

0.08

0.010

0.05

0.4 (0.1) 0.3 (0.1) 0.4 (0.1)

In addition, SEBM offers much higher AM efficiency (80 cm3 h
1) and greater flexibility for dealing with coarse powder (up to 180 μm; usually 40–90 μm) than selective laser melting (SLM) (typically 20–40 cm3 h
1 with smaller than 50 μm powder). Another characteristic of the SEBM process arises from the use of high substrate preheating temperatures (e.g. 730°C) and the preheating operation of each newly spread layer of powder during SEBM. This ensures that, in the case of Ti-6Al-4V, the powder bed temperature is maintained at temperatures above the minimum stress relief temperature (480°C) required for Ti-6Al-4V [5]. Consequently, SEBMfabricated Ti-6Al-4V does not normally require post-AM heat treatments while still possessing high tensile elongation (e.g., 15%) and good strength [7]. More than 3000 Ti-6Al-4V implants have been manufactured by the authors using SEBM over the last 5 years. This chapter discusses the design and manufacture of three SEBM-manufactured ELI Ti-6Al-4V implants that have been used in patients.

3.3.2

Feedstock material and the SEBM manufacturing process

ELI Ti-6Al-4V spherical powder was used for the manufacture of all Ti-6Al-4V implants described in this chapter. The composition of the powder, listed in Table 3.3.2, meets the requirements of ASTM F3001–14 [8]. The borderline oxygen level between ELI Ti-6Al-4V powder and non-ELI Ti-6Al-4V powder is 0.13%. Research has shown that ELI Ti-6Al-4V powder with an oxygen content of 0.08% can be reused up to four times before the oxygen content increases to 0.13% [9]. Table 3.3.2

Compositions of ELI Ti-6Al-4V spherical powder Composition (wt%)

H

C

O

N

Fe

Al

V

Ti

0.003

100 parts in eight different categories have been implanted into patients’ bodies in China in the as-built condition without surface treatment. All these SEBM-fabricated Ti-6Al-4V implants have performed satisfactorily in patients’ bodies. By 2040, >20% of the population in China will be over 65 years old. Among them, those who are over 80 will number >74 million. Consequently, there is a significant demand for various types of bone implants in China for aged people alone. Similar to the developments we have seen in Europe and other countries over the last decade, it is predicted that SEBM-fabricated Ti-6Al-4V orthopedic implants will find increasing applications in the near future in China.

Acknowledgments This project was partly supported by the Ministry of Science and Technology of China through grant 2016YFB1101400. The authors acknowledge the fruitful collaborations with Professor Ma Qian of RMIT University, Melbourne, Australia.

References [1] Arcam History, http://www.arcam.com/company/about-arcam/history/ (accessed 24.01.17). ® [2] P. Hollowaty, Arcam CAD to Metal , http://www.slideshare.net/PaulHollowaty/arcamebmelectron-beam-melting, 2016. Accessed 2 January 2017. [3] L. Zhongjun, 3D printing reshapes orthopedic surgery, China Today (2016). http://www. chinatoday.com.cn/english/report/2016-10/11/content_728724.htm. Accessed 2 April 2017. [4] L. Youle, The first successful implantation of 3D artificial cervical spine by PKU Third Hospital, http://english.pku.edu.cn/news_events/news/focus/5287.htm, 2017. Accessed 2 April 2017. [5] I. Polmear, D.H. StJohn, J.F. Nie, M. Qian, Light Alloys: Metallurgy of Light Metals, fifth ed., Butterworth-Heinemann, 2017. [6] ISO5832-3:2016, Implants of surgery—Metallic materials—Part 3: Wrought titanium 6-aluminium 4-vanadium alloy. [7] M. Qian, W. Xu, M. Brandt, H.P. Tang, Additive manufacturing and post-processing of Ti-6Al-4V for superior mechanical properties, MRS Bull. 41 (2016) 775–783. [8] F3001-14 A, Standard Specification for Additive Manufacturing Titanium-6 Aluminum-4 Vanadium ELI (Extra Low Interstitial) With Powder Bed Fusion, ASTM, West Conshohocken, PA, 2014. [9] H.P. Tang, M. Qian, N. Liu, X.Z. Zhang, G.Y. Yang, Effect of powder reuse times on additive manufacturing of Ti-6Al-4V by selective electron beam melting, JOM 67 (3) (2015) 555–563. [10] TIMETAL, Properties and processing of TIMETAL 6--4, http://www.timet.com/images/ document/technicalmanuals/TIMETAL 6--4_Properties.pdf, 1998. Accessed 1 December 2017.

Ti-6Al-4V orthopedic implants made by selective electron beam melting

249

[11] H.P. Tang, J. Wang, C.N. Song, N. Liu, L. Jia, J. Elambasseril, M. Qian, Microstructure, mechanical properties and flatness of SEBM Ti-6Al-4V sheet in as-built and hot isostatically pressed conditions, JOM 69 (3) (2017). [12] Y.Y. Sun, S. Gulizia, C.H. Oh, D. Fraser, M. Leary, Y.F. Yang, M. Qian, The influence of as-built surface conditions on mechanical properties of Ti-6Al-4V additively manufactured by selective electron beam melting, JOM 68 (3) (2016) 791–798.

3D-printed titanium alloys for orthopedic applications

3.4

A. Kumar, R.D.K. Misra University of Texas at El Paso, El Paso, TX, United States

3.4.1

Introduction

The skeletal disorders induced by injury, trauma, and diseases such as tumors or osteoporosis can be treated with temporary or permanent implants [1]. The critical size defects, also called long-bone defects, are difficult to repair even in the presence of mechanical support [2–4]. Therefore, filling the defect with bioactive material is required to accelerate bone regeneration and to solve the problem of nonunion of critical size defects. In this context, chemical and mechanical properties are relevant parameters for long-term performance of an implant in physiological conditions. To select a biomaterial for orthopedic applications, it is desirable to know the properties of bone. Human bone is characterized by a porous inner part with a compressive strength of 7–10 MPa, referred to as cancellous bone. This cancellous bone is surrounded by a dense cortical bone of compressive strength of 170–193 MPa [5,6]. Thus, to design bone analogue material, it is desirable to mimic properties of bone such as architecture, porosity, structure, and composition. However, no single material exhibits these properties. For instance, porous biomaterial provides the necessary space for vascularization and an environment for cells to grow in the porous scaffold. But porous scaffolds have inferior mechanical properties. In contrast, dense biomaterials are stronger and capable of withstanding the applied load but fail in providing the natural environment for cells to form bone tissue. In the event an implant is weaker than the bone, the defect can be stabilized with artificial devices such as metal plates, screws, and nails. In the case of implant materials with an elastic modulus higher than that of the bone, bone resorption may occur because of the stress shielding effect [7]. The stress shielding effect can be minimized using a porous metallic structure [8,9]. In this context, the metallic implants with lower strength and modulus than bone because of porosity may fail during loading because of excessive load transfer to the implant [10–12]. In this case, the pore architecture needs to be carefully tuned to increase the load-bearing capabilities of the implants. It is important to remember that the mechanical properties of an implant depend on apparent density and pore architecture of the implant [13,14]. As an alternative to the conventional scaffold fabrication techniques such as salt leaching and freeze casting, additive manufacturing (AM) has found diverse applications in the biomedical engineering field because of its effectiveness in the fabrication Titanium in Medical and Dental Applications. https://doi.org/10.1016/B978-0-12-812456-7.00012-3 © 2018 Elsevier Inc. All rights reserved.

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of patient-specific implants and control on pore architecture, pore-size distribution, and mechanical properties [15,16]. Depending on the application and materials, a number of 3D printing methods are now available [17–19]. Based on the ASTM standard F2792-12a, 3D printing methods can be classified into seven categories: (i) binder jetting, (ii) direct energy deposition, (iii) material extrusion, (iv) material jetting, (v) powder bed fusion, (vi) sheet lamination, and (vii) vat photopolymerization. Among direct energy deposition methods, electron beam melting (EBM), selective laser sintering (SLS), and laser-engineered net shaping (LENS) are specifically used for the fabrication of metallic scaffolds; they are covered in this chapter. Besides mechanical properties, cell-material interaction plays a major role in dictating the success of the implant [20,21]. The cell-material interaction can be modulated using surface chemistry and texture [22–28]. This includes the creation of a nanotextured surface and coating of the implant surface with extracellular matrix-like materials. In the sections below, we discuss the methods and in vitro and in vivo results to illustrate the efficacy of these methods in enhancing cell proliferation and osseointegration.

3.4.2

Metallic biomaterials

In the context of load-bearing applications among various mechanical and physical properties, elastic modulus has been extensively studied (Fig. 3.4.1) [29]. This is because of the higher elastic modulus of implant results in the stress shielding effect because of higher load transfer to the implant in relation to the bone [30]. This triggers Causes of implant failure

Wear debris generation/corrosion

Inflammation

Fibrous encapsulation

Particulate debris accumulates in the human tissue

Low fatigue strength/low fracture toughness

Mismatch in modulus of elasticity between implant and host bone

Fracture of implants due to mechanical failure

Stress shielding and bone resorption

Hypersensitivity, tissue toxicity, carcinoginity Revision surgery

Co-Cr alloys show high wear resistance

Ti-alloys are free of toxic elements and are biocompatible

Ti-alloys exhibit high specific strength

B-type Ti-alloys exhibit low elastic modulus

Fig. 3.4.1 Possible reasons for failure of the implant and the role of biomaterial properties on implant performance [29].

3D-printed titanium alloys for orthopedic applications

253

the remodeling of the bone in the absence of sufficient loading, which results in bone resorption due to excessive osteoclast activity in the stress-shielded region of the bone [31]. Other causes of implant failure are low fatigue strength and poor fracture toughness. The poor fracture toughness leads to implant fracture during loading and low fatigue strength shortens the life of the implants. Other factors responsible for implant failure are generations of wear and corrosion debris in the vicinity of articulating implants, which may cause local and systemic cytotoxicity [29]. The release of metallic ions due to wear and corrosion may lead to hypersensitivity. Additionally, the carcinogenic effect of this metallic debris cannot be ignored. Another possible cause of rejection of metallic implants is inflammation. Also, the formation of a fibrous capsule around the bioinert metallic implants (for instance, titanium) is responsible for nonbonding or integration of the implant with the host bone. This results in the painful failure of the implant. In this context, the coating of the implant with bioactive material is a possible solution to reduce the inflammation as well as to avoid fibrous encapsulation. To overcome the above-mentioned problems and to obtain a solution, metallic biomaterials are mainly used in the reconstruction of damaged hard tissue. The most popular metallic biomaterials for the load-bearing application are Ti alloys, Co-Cr alloys, and austenitic stainless steels [32]. Among these biomaterials, Ti alloys exhibit the highest specific strength, corrosion resistance, and biocompatibility [32–35]. As compared to Ti alloys and austenitic stainless steels, Co-Cr alloys show greater fatigue strength and wear resistance [32]. Stainless steels exhibit higher ductility and, therefore, different devices can be fabricated easily [32,36]. Among these metallic biomaterials, Co-Cr alloys and Ti alloys have a maximum and minimum elastic modulus, respectively [32,33,37]. In addition to this, the alloying element nickel (Ni) used up to 20% in austenitic stainless steel to stabilize the austenitic phase is widely recognized as a high-risk element and can cause inflammatory hypersensitivity reactions [32,38].

3.4.3

Titanium alloys

Ti alloys have been widely used for bone reconstruction surgery, including total hip replacement (THR), total knee arthroplasty (TKA), dynamic compression plate (DCP), and lumbar fusion and fixation [39–44]. Titanium alloys exhibit high corrosion resistance, a relatively lower elastic modulus, and biocompatibility [32–35]. As mentioned above, a biomaterial with an elastic modulus similar to bone can minimize stress shielding. Titanium alloys have a lower elastic modulus than Co-Cr alloys and stainless steels [37]. In this context, the Ti alloy Ti-6Al-4V with (α + β)-phase is characterized by an elastic modulus of 110 GPa. This elastic modulus is significantly lower than the elastic modulus of Co-Cr alloys (210 GPa) and stainless steels (180 GPa). Depending on the phase assemblage, Ti alloys can be divided into three categories: α-type, β-type, and (α + β)-type. The elastic modulus of β-Ti alloys is lower than α- and (α + β)-Ti alloys. Therefore, β-Ti alloys are suitable for load-bearing applications to reduce bone resorption caused by the stress shielding

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Young's modulus (GPa)

250 200 SUS 316L CoCrMo 150 Cp-Ti Ti-64ELI Ti-13Zr-13Nb TNZTO TFNZ

100 TNTZ

Ti-8Mo

50

Bone

Z11

0 0

0.2

0.4 0.6 0.8 1.0 1.2 Elastic admissible strain, e (%)

1.4

1.6

Fig. 3.4.2 Elastic admissible strain versus elastic modulus (Young’s modulus) for bone and different biomaterials [29].

effect [31]. In the past, the developed β-Ti alloys for biomedical applications were Ti-16Nb-10Hf [45], Ti-29Nb-13Ta-4.6Zr [46], Ti-13Nb-13Zr [47], Ti-12Mo-6Zr2Fe [48], Ti-15Mo [49], Ti-15Mo-5Zr-3Al [50], and Ti-35.3Nb-5.1Ta-7.1Zr [51]. Considering the use of metallic biomaterials for load-bearing orthopedic applications, the yield stress-to-modulus ratio, referred to as elastic admissible strain, is an important parameter (Fig. 3.4.2) [29]. High strength and low elastic modulus are the most important properties after biocompatibility and corrosion resistance for designing an implant. Thus, a material with higher elastic admissible strain is most desirable for orthopedic applications [52].

3.4.4

Toxicological effect of titanium alloys

Although biomedical grade metals perform reasonably well in corrosive environments, there are reports of ion exchange at the interface of the implant and the physiological fluid. The amount of ionic current depends on the corrosion resistance property of the implant material in the physiological fluid [53]. Corrosion resistance is strongly affected by environmental conditions such as pH and chloride ion concentration. Other factors that influence the corrosion resistance property are preexisting cracks, surface abrasions, and coating properties. Additionally, galvanic effect due to the presence of dissimilar phases as well as localized cellular activity (cells growing on the surface of the sample), crevice, and pitting corrosion leads to the production of corrosion products [54–56]. Thus, the corrosion products generated due to these processes may affect cellular and tissue function because of local and systemic toxicity. Besides the aforementioned factors, the elements released in the physiological environment due to degradation of the biomaterial may lead to severe long-term adverse pathological effects [57,58]. As mentioned in Section 3.4.3, Ti alloys often

3D-printed titanium alloys for orthopedic applications

255

contain Nb, V, Al, Fe, Ta, Mo, and various other elements. Considering the tissue reaction in the presence of these elements, V, Al, and Nb (or Ti, Zr, and Ta) can be categorized as sterile [59]. From the biological point of view, in the case of long-term implantation, Ni has been studied for its allergenic and carcinogenic effects; it is known for severe sensitization [60,61]. Aluminum was associated with Alzheimer’s disease, a severe neurological disorder, as well as in osteomalacia [62]. Therefore, efforts were made to develop V- and Al-free Ti alloys such as the Ti-13Nb-13Zr alloy to improve bone biocompatibility [29,63].

3.4.5

3D printing of titanium alloy scaffolds

Conventionally, freeze drying [64,65], solvent casting in combination with particulate leaching [66], fiber bonding, and gas foaming [67] are used for scaffold fabrication. Although these methods are capable of producing 3D scaffolds, it is extremely difficult to control pore size, shape, distribution, and pore interconnectivity [68]. In contrast, 3D printing methods are capable of producing 3D scaffolds with highly intricate and patient-specific designs. Also, it is possible to control pore size (to a certain extent), shape, distribution, and interconnectivity in the 3D printing of scaffolds [26,27,69]. EBM, SLS, and LENS are commonly used to make metal implants. In brief, an EBM system comprises an electron gun, operating at an anode potential of 60 kV, that is scanned over a uniform layer of metal powder [70,71]. The scan is driven by a CAD design to selectively melt the metal particles. After printing, the second layer of powder is deposited from a gravity-driven feeder. In contrast to EBM, SLS utilizes a laser as a source of energy to melt the uniform layer of metal powder [72]. The laser-deposited energy leads to the melting of metal particles, followed by solidification. After one layer is printed, the next layer is deposited by spreading the metal powder from a feed platform using a dispensing roller. In the LENS method, a metal powder is sprayed at the focal point of a high-power laser to melt and deposit the powder [73]. In this, the CAD controls the scan to fabricate the desired model. Technically, the LENS process is similar to other direct energy deposition-based 3D printing methods. However, this method leads to the formation of a heat-affected zone in the metallurgically bonded region. Orthopedic devices of metallic biomaterials, such as stainless steels, Co-Cr alloys, and Ti alloys, can be successfully fabricated using 3D printing methods. However, an additional treatment is required to modify the surface to improve the cellular activity, in vivo. The details of different surface modification methods and their effects on the biocompatibility are discussed in the next section. Apart from this, to design scaffolds with properties close to the intramedullary bone regime, we need a metallic scaffold with pore architecture to bring the value of Ebone/Escaffold close to 1. In general, for a solid Ti-6Al-4V scaffold with Escaffold 110 GPa, the value of Ebone/Escaffold is 0.09 [74]. To increase the Ebone/Escaffold value close to 1, we require a mesh or foam structure. In addition to this, the higher surface area due to porous architecture helps in cell attachment and proliferation. The higher surface area further provides the space for vascularization that ensures the supply of blood to the cells growing deep inside the scaffold.

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To study the effect of pore cell architecture on mechanical properties, Ti-6Al-4V mesh structures with different elements, such as cubic or rhombic dodecahedrons, were fabricated using the EBM method [75]. Mechanical testing revealed a compressive strength and elastic modulus of 10–300 MPa and 0.5–15 GPa, respectively. In a different study, a solid prototype of Ti-24Nb-4Zr-7.9Sn was printed using EBM. The microindentation on the printed samples revealed a hardness of 2.5 GPa. A similar value of hardness (2.3 GPa) was found in the case of a solid component fabricated by SLS [76]. In a different study, SLS was used to fabricate Ti-6Al-4V alloy scaffolds with low interstitial content of four different combinations of strut thickness and porosities (120, 500; 170, 450; 170, 500; and 230, 500 μm) with a dodecahedron unit cell to study the effect of porosity on mechanical properties. Cylindrical specimens of dimensions 10 mm  15 mm were fabricated using SLS. Mechanical tests indicated a variation in the yield stress from 16 to 92 MPa and in the maximum stress from 19 to 117 MPa with an inverse relation between strength and porosity, that is, the lowest strength in the case of the highest porosity [77]. It is important to remember that optimization among mutually opposing properties, such as high porosity and high strength with high-energy absorption capabilities, is required to design a patient-specific implantable device. These requirements limit the choice of biomaterials for load bearing orthopedic applications. Therefore, a design specific approach is needed to overcome these limitations. In this context, 3D printed gradient Ti-6Al-4V mesh structures were found to be promising [78]. The compressive strength measurement revealed a structure-dependent stress distribution during loading. The results indicated a deformation behavior of the entire graded structure that can be considered as a weighted average of each uniform mesh element.

3.4.6

Biocompatibility of titanium alloys

The biocompatibility of an implant is defined by its performance in a physiological environment, which is assessed by its local and systemic toxicity, genotoxicity, and effect on immune response. Therefore, biomaterials must be characterized for these properties to minimize the risk of failure of the implant as well as the adverse effect on the patient’s health. In the context of bone tissue engineering applications, osseointegration of the implant with the host bone is required for long-term performance. After implantation, a monolayer of protein is adsorbed on the surface of the implant within seconds of contact with blood (Fig. 3.4.3). The adsorption of proteins and their conformation is governed by the surface properties of the implant [21,79–81]. Furthermore, these proteins and their conformational state subsequently affect cell-material interaction. Although the exact mechanisms are not well understood, it is reported that the chemical and physical properties play a decisive role in the cell-material interaction. The implant surface can affect osteogenic gene expression through modulation of cellular function under physiological conditions [82–84]. There are various factors that can affect cell-material

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Cell with integrin receptors

Proteins

Substrate

(A)

(B) Fig. 3.4.3 Cell-material interaction involves the adsorption of proteins immediately after first contact of the implant with bodily fluids (A). Adsorbed protein can efficiently improve the cell-material interaction through the activation of integrin receptors (B).

interaction, and therefore the biocompatibility of the implant. These factors include surface chemistry as well as both physical and mechanical properties.

3.4.6.1 Effect of surface chemistry Surface chemistry is an important factor in cell-material interaction. Surface chemistry determines the surface energy and thus the wettability of the surface. It has been observed that protein adsorption and cell adhesion is promoted by the hydrophilic surface. In general, a titanium surface is negatively charged because of adsorbed anions, such as OH and F , from the solution. Given that the cell membrane is also negatively charged, the presence of a positively charged layer of protein present on the titanium surface can enable initial cell attachment [85]. In this regard, surface topography can positively contribute to cell attachment because of a change in surface

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chemistry. An important fact is that the above only affects the surface without changing the bulk properties. In contrast, a few studies reported that anions present in the oxide layer topography are responsible for cell attachment [86]. Thus, the composition of the oxide layer is a more effective parameter than the topography of the oxide layer. In this context, different ions have different effects on cell response. For instance, F stimulates the initial cell attachment and PO43 encourages proliferation [86]. Although nanotopography can be used to modulate osteoblast functions, immobilization of different bioactive molecules on the biomaterial surface can efficiently enhance cell adhesion and proliferation, in vitro [87]. In this context, peptide fragments from the cell-extracellular matrix (ECM) proteins such as fibronectin and arginine-glycine-aspartic acid, were found to be very effective in stimulating the cell-material interaction and cell adhesion through activation of integrin receptors [88,89]. In another example, sericin, a water-soluble protein, immobilized on the Ti alloy surface was found to be very effective in modulating osteoblast differentiation [90,91]. Furthermore, macrophage cells and osteoblast-macrophages were cocultured on a sericin-immobilized titanium surface. Results revealed insignificant amounts of proinflammatory cytokines such as TNF-α and IL-1β [92]. Because Ti alloys are bioinert in nature, a surface treatment is required to activate the cellular activity on the biomaterial surface and to improve the biocompatibility of the implant. Nanotopography was observed to be very effective in modulating osteoblast functions, which can be achieved by acid etching or oxidation. The coating of a bioactive material, such as calcium phosphate, can also be used to improve the bioactivity of a Ti alloy surface. Moreover, immobilization of a surface with a protein similar to ECM can efficiently improve cell-material interaction through the activation of integrin receptors. It has been reported that the cell phenotype, functions, and interactions with an ECM depend on the physiognomies of ECM [93–97]. In this context, a tissueengineered environment similar to natural ECM can be used to generate a favorable biological response [98]. Considering this aspect, a decellularized extracellular matrix (dECM) was utilized to provide an optimal environment for cells similar to native tissue ECM, because synthetic biomaterials are incapable of reproducing the characteristics of the ECM [93]. In this regard, efforts were made to utilize the dECM of adipose and liver with bioink during 3D bioprinting to design the tissue construct for enhanced biocompatibility and biofunctionality of the biomaterial [93–97]. Recently, an in vitro cell culture approach was employed to produce dECM directly on the scaffold surface [28,99,100] and to provide the essential cues for cell function [101,102]. Kumar et al. used [100] in vitro method to grow the osteoblasts over the Ti-6Al-4V surface, followed by osteogenic differentiation. The differentiated cells were subjected to multiple freeze-thaw cycles to decellularize the extracellular matrix (Fig. 3.4.4) [100]. dECM-coated Ti-6Al-4V samples were used as an ideal substrate for the cell culture experiments. Cytocompatibility assessment revealed a higher cell adhesion and proliferation on dECM/Ti-6Al-4V scaffolds as compared to bare Ti-6Al-4V scaffolds. The enhanced cellular functionality can be related to the extracellular matrix peptides and proteins, which were present in the dECM [58,59]. A higher density of the actin filament bundle was found on the dECM-coated scaffolds than on the bare Ti-6Al-V (Fig. 3.4.5) [100]. In this context, it is important to mention that vinculin and actin are

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Cell culture and differentiation of cells

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Decellularized scaffold

Three dimensional morphology of cells growing on dECM/scaffold

Extracellular matrix ornamented scaffold (dECM/scaffold)

Cell culture on dECM/scaffold

Ti-6Al-4V

Fig. 3.4.4 Process of decellularization to produce the osteoblasts extracellular matrix on the Ti-6Al-4V scaffolds surface [100].

100 µm

(B)

(C)

100 µm

(D)

dECM/Ti-6Al-4V

(A)

100 µm

100 µm

Fig. 3.4.5 Immunochemistry images showing the bundle of actin stress fibers after seven days of osteoblast culture on the pore region (marked with white broken lines) and struts of (A, B) Ti-6Al-4V and (C, D) dECM ornamented T-6Al-4V scaffolds [100].

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responsible for the transduction of mechanical stimuli to the cytoskeleton and cell motility, respectively [103–106]. The expression level of these proteins is indicative of the degree of the cell-to-cell and cell-to-material (matrix) interactions.

3.4.6.2 Effect of surface topography Although Ti alloys exhibit superior biological properties versus Co-Cr alloys and stainless steels, a bioinert surface restricts the osteogenic activity on Ti alloy surfaces [107,108]. Therefore, to promote osseointegration, various surface modification methods have been explored such as grit blasting, calcium phosphate coating, acid etching, and anodic oxidation [109–111]. The effect of titania nanotube diameter on the water contact angle was studied by Hao et al. [112]. A significant effect of nanotube diameter on the contact angle was noted such that nanotubes of larger diameter had smaller contact angles (Fig. 3.4.6) [112]. A sharp decrease in the contact angle was observed for nanotubes with a diameter in the range of 50–70 nm, with no change in contact angle on further increase in nanotube diameter. This decrease in contact angle was related to an increase in surface energy with increasing nanotube diameter [112]. Surface topography can change the surface energy by changing the surface roughness and, therefore, can affect the cell-material interaction. In this context, a higher extension of filopodia on the nanotextured surface was found [113]. Thus, a significant effort has been made to create nanosized features on the biomaterial surface. The electrochemical anodic oxidation method was observed to be promising in nucleating a controllable and uniform nanopattern of titania nanotubes on the titanium surface

70 nm

50 nm

70 Surface contact angle (*)

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60

*

50

*

40 30

*#

20

*#

10 0

200 nm

(A)

Ti

Ti2448

30 nm

50 nm

70 nm

90 nm

(B)

Fig. 3.4.6 SEM micrograph showing the self-aligned titania nanotubes of different diameters (A) and their effect on the water contact angle (B). Statistical analysis shows the significant difference at P < 0.05 as compared with pure Ti and Ti2448 groups (marked with *). # represents the statistically significant difference P < 0.05 as compared with the 30 and 50 nm groups [112].

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[114]. The titania nanotubes can be obtained by anodic oxidation of titanium in an electrolyte. A constant potential is applied between a titanium working electrode and a counter electrode (platinum) in an electrochemical cell. The titania nanotube features, such as diameter, wall thickness, and length, can be controlled by controlling anodization parameters such as electrolyte concentration and pH, applied voltage, current density, and anodization time [115]. Even in the nanoscale regime, different sizes of nanotubes can affect cell adhesion in a different manner. In one study, titania nanotubes of 15 nm diameter exhibited optimal adhesion and differentiation of cells [116]. In another study, a higher focal adhesion was noted on nanotubes of 30 nm diameter as compared to larger nanotubes. In a similar work, higher osteoblast cell adhesion was noted on titania nanotubes of 30 nm diameter. In addition to this, an enhanced cell elongation and ALP (alkaline phosphatase) activity were observed on nanotubes of diameter between 70–100 nm [117]. In a different study, a strong effect of the nanotube diameter on the osteoblast cell morphology was observed (Fig. 3.4.7) [112]. After a 3 h incubation of osteoblast (MG-63) on nanotubes, a higher spreading of cells was noted on 30 nm diameter nanotubes, which decreased with increasing nanotube diameter [112]. A similar result was found after 24 h incubation. A higher spreading of cytoskeletal actin was found on the surface with a smaller nanotube diameter (30 nm) than the surface with large diameter nanotubes (50–90 nm) [112]. It is important to mention that the water contact angle decreased with increasing nanotube diameter with the highest contact angle being measured on

Ti

Ti2448

30 nm

50 nm

70 nm

90 nm

3h

24 h

50 µm

Fig. 3.4.7 Immunofluorescent micrographs showing the morphology of the osteoblast cells (MG-63) cultured on commercially pure Ti Ti2448 as well as titania nanotubes of diameters 30, 50, 70, and 90 nm, after 3 and 24 h incubation [112].

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nanotubes with a smaller diameter (30 nm). Previous studies emphasized the importance of a hydrophilic surface with superior cell response on the more hydrophilic surface [118–120]. On the basis of these arguments, one can consider that topography is a more dominant factor than surface wettability. Therefore, cell adhesion decreases with increasing nanotube diameter. However, larger-diameter nanotubes promoted cell differentiation [115].

3.4.7

Vascularization of 3D scaffolds with designed porous architecture

One of the main challenges in tissue engineering is the vascularization of the 3D construct. The absence of proper oxygen transport in scaffolds leads to necrotic cell death in the core of scaffolds. Therefore, pore architecture of a scaffold determines vascularization and oxygen diffusion. The formation of new blood vessels is described by the angiogenesis (Fig. 3.4.8A) [121,122], which is affected by the scaffold architecture such as pore size, shape, and interconnectivity [123]. Given that, bone is characterized by a highly vascularized tissue [124–126]. Therefore, the success of a bone-analogue 3D scaffold depends on the efficacy of the scaffold to provide oxygen and nutrient supply without interruption. In this context, the limiting parameter for oxygen supply from blood vessels to neighboring cells through diffusion depends on the diffusion length of oxygen, equal to 150–200 μm (Fig. 3.4.8B) [124,127,128]. As mentioned before, hypoxia introduced by the limited supply of oxygen and nutrients leads to cell necrosis [129]. Therefore, pores of 100–200 μm size are considered suitable for the angiogenesis [130–132] and for the oxygen diffusion, the maximum distance between the capillaries should not exceed 150–200 μm [133,134]. Moreover, pore size in the range of 300–400 μm was found suitable for both angiogenesis and bone ingrowth [135,136]. Vascularization is a slow process [137] and takes a few weeks to grow a capillary of a few millimeters [124]. Therefore, a pretreatment of scaffolds with vascular growth factors is often required to stimulate the blood capillary in-growth in bigger scaffolds [124]. The most common growth factor for the angiogenesis is VEGF (vascular endothelial growth factor) [138], which helps in faster healing of bone defects due to the formation of a vascular network [139]. For instance, as a result of vascularization, a threefold increase in new bone formation was noted in Balb/c mice after 28 days due to an improved supply of oxygen and nutrients to osteoprogenitor cells [140,141]. In a recent study, Lv et al. [142] utilized the growth factor-doped fibrin glue to coat the EBM-fabricated titanium scaffold to promote vascularization and osseointegration. Fibrin serves as a carrier for the BMP-2 and VEGF and allows a gradual release of these proteins from the porous structures because of slow degradation of fibrin glue with time. The slow degradation of fibrin glue allows the localized delivery of BMP-2 and VEGF and, thus, facilitates the angiogenesis and osteogenesis in the interior of the titanium scaffold. A four-week study associated with the angiogenesis and

Diffusion & Transport into the blood

Medium size vessel Oxygen

O2

CO2

Nutrients

Carbon dioxide Waste products

Drugs

3D-printed titanium alloys for orthopedic applications

Diffusion & Transport into the tissue

Large vessel

Capillary vessels Maximum distance 200 µm

(A)

(B)

Fig. 3.4.8 Schematic showing (A) the angiogenesis in a 3D porous scaffold [122], (B) relation of diffusion length of oxygen with intercapillary distance [166].

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osteogenesis in the critical size defect in the rabbit medial femoral condyle revealed a significant increase in the angiogenesis. However, no improvement in the osteogenesis was observed [142]. Other studies [143–145] also reported the absence of new bone formation in the presence of both BMP-2 and VEGF, which is possibly due to differentiation of bone marrow stem cells into an endothelial lineage in the presence of VEGF and, thus, limiting the number of cells for osteogenesis in the defect area [146]. Therefore, an optimum ratio of BMP-2/VEGF can be useful in promoting the osteogenesis differentiation of stem cells in addition to angiogenesis via the differentiation of stem cells into endothelial cells. In a different work, Matena et al. [147] used the proangiogenic factors to promote vascularization in porous titanium scaffolds fabricated by SLM. To accomplish this, a titanium scaffold was coated with polycaprolactone, followed by functionalization of coating with proangiogenic factors such as VEGF and high mobility group box 1 (HMGB1). Results showed a higher chemotactic potential on endothelial cells in the presence of HMGB1 as compared to VEGF, in vitro. In summary, angiogenesis is a key requirement for the successful use of 3D scaffolds in the reconstruction of bone defects because vascularization allows the growth of blood capillaries in the inner part of the scaffold and thus facilitates the supply of oxygen and nutrient. Although titanium implants fail to provide a suitable environment for vascularization, a chemotactic potential such as VEGF can be used to promote the growth of blood capillaries. However, an optimal ratio of BMP-2/VEGF can be useful in promoting both osteogenesis and angiogenesis.

3.4.8

Antibacterial effect of titanium alloys

Apart from poor osseointegration and aseptic loosening, bacterial infection is a major problem [148,149]. With the increase in surgical reconstruction, there is a corresponding increase in bacterial infection. Implant-related infection is a serious problem when a bacterial colony gains resistance to the natural host defense system. In extreme cases, when antibiotic treatment fails, surgical intervention is required to remove the infected part [150]. In this regard, significant effort has been made in recent years toward the development of implants with anti-infective surfaces [151]. These surfaces are expected to prevent bacterial colonization and thereby decrease the dependency on antibiotics. Sources of bacteria are not only limited to the implant itself but also to the operating room atmosphere, the medical professional’s hygiene, and the patient’s own skin and bodily interior [150,152,153]. Staphylococci-related infection is the major cause of implant-related infection, which is responsible for at least four out of five infections [154,155]. After implantation, there is a race to the surface between bacteria and cells. If the race is won by bacteria, then cells will not be able to grow over the surface in the presence of colonized bacteria [156]. The colonization of bacteria is followed by biofilm formation, which protects bacteria from antimicrobial agents and the host immune system. Bacteria protected in the biofilm is several orders of magnitude more resistant to antibacterial agents than bacteria

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in the planktonic state [157]. The initial 6 h after implantation is considered as the “decisive period” when the implant is highly susceptible to infection. Therefore, the protection of the implant from bacterial colonization is required for the long-term success of the implant [158]. There are two approaches to overcome the bacterial infection-surface topography and an implant surface loaded with an antimicrobial agent. Considering the size of cells and bacteria, osteoblast is significantly bigger than the bacterial. Also, the stiffness of both cell and bacteria is significantly different. Therefore, a surface topography with modified stiffness to only support the osteoblast interaction with the implant surface can be an effective solution in limiting the bacterial infection of the implant [159]. In another approach, an antibacterial agent such as silver (Ag) can be introduced onto the surface of the titanium implant to suppress bacterial colonization. Ag is known for its bactericidal effect against a broad spectrum of bacteria [160,161]. Although Ag is proven for its bactericidal properties, its therapeutic window is small, above which it is cytotoxic [162,163]. For loading and sustained release of antimicrobial agents, a carrier material is required. In this context, titania nanotubes can be used as a reservoir to store the antibacterial agents such as silver [160,164,165]. In a study by Mei et al. [165], the bactericidal effect of silver-implanted titania nanotubes was explored (Fig. 3.4.9). in vitro and in vivo studies revealed a significant bactericidal effect of titania nanotube samples plasma-implanted with silver at a voltage of 0.5 kV (NT-Ag-0.5) and 1 kV (NT-Ag-1.0). In a different work, a titanium-copper alloy was prepared to prevent bacterial adhesion and colonization. Copper is expected to protect the implant from bacterial infection without deteriorating the mechanical and corrosion-resistant properties.

3.4.9

Conclusions

Trauma or age-related diseases are responsible for skeletal disorders that severely affect physiological functions. The bone tissue function can be restored using a suitable implant material. Among various implant materials, titanium alloys are highly appropriate for bone replacement because of their high strength, low elastic modulus, and high corrosion-resistance properties. However, a bioinert surface limits the wider application of titanium alloys and surface modification is required to improve the biocompatibility of the implant material. Apart from this, the presence of toxic elements such as Al and V in a commonly used titanium alloy (Ti-6Al4V) raises the question of long-term performance of implant material in physiological conditions. Therefore, titanium alloys with different alloying elements such as Ti-13Nb-13Zr have been developed to improve the long-term performance of titanium alloys. Furthermore, the titanium alloy surface can be modified to grow the titania nanotubes to promote cell attachment and proliferation. These nanotubes can also be used as a reservoir to store the antimicrobial agent such as silver to prevent bacterial colonization on the implant after implantation.

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PT

PT

Aa

Pg NT

NT

Aa

Pg NT-Ag-0.5

Pg

NT-Ag-0.5

Aa NT-Ag-1.0

NT-Ag-1.0

Pg

Aa

Fig. 3.4.9 SEM micrographs of bacteria, Porphyromonas gingivalis (PG) and Actinobacillus actinomycetemcomitans (Aa), on the sample surface after 1 day of incubation, showing the lower bacterial cell viability on NT-Ag-0.5 and NT-Ag-1.0 than NT and PT. PT denotes the untreated pure Ti and NT after anodization (to grow the nanotubes) at 20 V for 30 min [165].

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[131] P. Carmeliet, Mechanisms of angiogenesis and arteriogenesis, Nat. Med. 6 (4) (2000) 389–396. [132] M.W. Laschke, Y. Harder, M. Amon, I. Martin, J. Farhadi, A. Ring, N. Torio-Padron, R. Schramm, M. R€ucker, D. Junker, J.M. H€aufel, C. Carvalho, M. Heberer, G. Germann, B. Vollmar, M.D. Menger, Angiogenesis in tissue engineering: breathing life into constructed tissue substitutes, Tissue Eng. 12 (2006) 2093–2104. [133] P. Carmeliet, R.K. Jain, Angiogenesis in cancer and other diseases, Nature 407 (6801) (2000) 249–257. [134] R.Y. Kannan, H.J. Salacinski, K. Sales, P. Butler, A.M. Seifalian, The roles of tissue engineering and vascularisation in the development of micro-vascular networks: a review, Biomaterials 26 (14) (2005) 1857–1875. [135] Y. Kuboki, Q. Jin, H. Takita, Geometry of carriers controlling phenotypic expression in BMP-induced osteogenesis and chondrogenesis, J. Bone Joint Surg. 83 (1 suppl 2) (2001) S105–S115. [136] E. Tsuruga, H. Takita, H. Itoh, Y. Wakisaka, Y. Kuboki, Pore size of porous hydroxyapatite as the cell-substratum controls BMP-induced osteogenesis, J. Biochem. 121 (2) (1997) 317–324. [137] E.R. Clark, E.L. Clark, Microscopic observations on the growth of blood capillaries in the living mammal, Am. J. Anat. 64 (2) (1939) 251–301. [138] C.R. Anderson, A.M. Ponce, R.J. Price, Immunohistochemical identification of an extracellular matrix scaffold that microguides capillary sprouting in vivo, J. Histochem. Cytochem. 52 (8) (2004) 1063–1072. [139] S. Ochman, S. Frey, M.J. Raschke, J.N. Deventer, R.H. Meffert, Local application of VEGF compensates callus deficiency after acute soft tissue trauma—results using a limbshortening distraction procedure in rabbit tibia, J. Orthop. Res. 29 (7) (2011) 1093–1098. [140] S. Bose, M. Roy, A. Bandyopadhyay, Recent advances in bone tissue engineering scaffolds, Trends Biotechnol. 30 (10) (2012) 546–554. [141] J. Trueta, K. Little, The vascular contribution to osteogenesis. II. Studies with the electron microscope, J. Bone Joint Surg. (Br.) 42 (2) (1960) 367–376. [142] J. Lv, P. Xiu, J. Tan, Z. Jia, H. Cai, Z. Liu, Enhanced angiogenesis and osteogenesis in critical bone defects by the controlled release of BMP-2 and VEGF: implantation of electron beam melting-fabricated porous Ti-6Al-4V scaffolds incorporating growth factordoped fibrin glue, Biomed. Mater. 10 (3) (2015)035013.   [143] A. Herna´ndez, R. Reyes, E. Sa´nchez, M. Rodrı´guez-Evora, A. Delgado, C. Evora, In vivo osteogenic response to different ratios of BMP-2 and VEGF released from a biodegradable porous system, J. Biomed. Mater. Res. A 100 (9) (2012) 2382–2391. [144] S. Young, Z.S. Patel, J.D. Kretlow, M.B. Murphy, P.M. Mountziaris, L.S. Baggett, H. Ueda, Y. Tabata, J.A. Jansen, M. Wong, Dose effect of dual delivery of vascular endothelial growth factor and bone morphogenetic protein-2 on bone regeneration in a rat critical-size defect model, Tissue Eng. A 15 (9) (2009) 2347–2362. [145] D.H.R. Kempen, L. Lu, A. Heijink, T.E. Hefferan, L.B. Creemers, A. Maran, M.J. Yaszemski, W.J.A. Dhert, Effect of local sequential VEGF and BMP-2 delivery on ectopic and orthotopic bone regeneration, Biomaterials 30 (14) (2009) 2816–2825. [146] H. Peng, V. Wright, A. Usas, B. Gearhart, H.-C. Shen, J. Cummins, J. Huard, Synergistic enhancement of bone formation and healing by stem cell-expressed VEGF and bone morphogenetic protein-4, J. Clin. Invest. 110 (6) (2002) 751–759. [147] J. Matena, S. Petersen, M. Gieseke, A. Kampmann, M. Teske, M. Beyerbach, H.M. Escobar, H. Haferkamp, N.-C. Gellrich, I. Nolte, SLM produced porous titanium

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Ti-6Al-4V lattice structures fabricated by electron beam melting for biomedical applications

3.5

S. Zhao, W.T. Hou, Q.S. Xu, S.J. Li, Y.L. Hao, R. Yang Institute of Metal Research, Chinese Academy of Sciences, Shenyang, China

3.5.1

Introduction

As an important implant material, titanium and its alloys have been widely used in the biomedical field for several decades. Among numerous titanium alloys, Ti-6Al-4V, which has α + β type microstructure, is a typical solid implant material for dental, orthopaedic, and maxillofacial applications. It benefits from its high strength as well as good resistance to corrosion and fatigue in physiological media. However, most titanium alloys show relatively high elastic modulus (about 90–115 GPa) while the modulus of cortical hard bone (the outer region of femoral and tibial bones) is much lower, commonly in the range of 16–20 GPa and that of cancellous or trabecular soft bone is about 1–4 GPa [1,2]. The mismatch between the mechanical properties (mainly elastic modulus) of the bone and metallic implants may hinder tissue cell in-growth, resulting in stress shielding and bone resorption that will ultimately lead to failure of the orthopaedic implants [3–6]. Recently, titanium alloy cellular structures have been successfully fabricated by additive manufacturing (AM) using the electron beam melting (EBM) method. Compared to other reported techniques for fabrication of cellular titanium and its alloys, AM technology has some unique advantages including its ability to create arbitrary complex three-dimensional structures, highly accurate and predictable porous structures, and extensive material selection [7–10]. At present, most studies on AM-produced porous titanium alloys have been focused on the (α + β)-type Ti-6Al-4V alloy, especially its fabrication process and mechanical properties [1,11,12]. Thus, this chapter will briefly introduce the relationship between morphology and the mechanical properties of EBM cellular structures with various types of space-filling unit structures by discussing the deformation behaviors, mechanical properties, and fatigue failure mechanisms of the Ti-6Al-4V porous structures. In addition, the biocompatibility of Ti-6Al-4V cellular structures for biomedical applications will also be discussed [13].

Titanium in Medical and Dental Applications. https://doi.org/10.1016/B978-0-12-812456-7.00013-5 © 2018 Elsevier Inc. All rights reserved.

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3.5.2

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Design of Ti-6Al-4V cellular structures

Different reticulated mesh structures as well as stochastic foam structures of the Ti-6Al-4V alloy were designed for evaluating the influence of porosity and lattice structures. For foam structures, the ligaments are stochastic, and the mechanical of whole bulk foams can be considered isotropous. They can be used to study the relationship between mechanical properties and porous size while for AM-produced porous biomaterials, their mechanical properties are highly dependent on the type of unit cell from which they are made [14–19].

3.5.2.1 Unit cell structures Besides the interconnected channels, the internal structure or microstructure of the cellular structures also has to be reasonably designed. For example, macroporous open-cellular structures with pore sizes ranging from 100 to 1000 μm, depending on the final application, have been achieved by EBM technology for fabrication of scaffolds in order to facilitate the ingress of seeded cells. Polyhedrons are defined as a three-dimensional (3D) object with bounded polygons whose side is shared by two polygons [20]. A polygon shows a two-dimensional shape composed of a cycle of line segments that can be divided into two categories: selfintersecting (edges cross other edges) and nonself-intersecting. For the shape, there are also two categories: convex and nonconvex. For scientific research, many polyhedrons are not suitable for fabricating and exploring because of their complexity. If we just focus on illustrating the underlying influence of the structure, there are two elimination criteria that should be avoided: a. Can only be repeated regularly in 3D space by the joining of vertices or edges. b. Complex structures having a large number of small faces.

Considering the above problems, 11 kinds of shapes are optional for basic unit cells, as shown in Fig. 3.5.1. According to the angles between the struts and loading direction, three typical structures were chosen: cubic, G7 (square pyramid), and rhombic dodecahedron. They were used as the unit cells and were repeated along three dimensions, filling the whole bulk.

3.5.3

Fabrication of Ti-6Al-4V cellular structures

As a typical representative of additive (layer) manufacturing method, the EBM fabrication, i.e., using electron beams as point source heating technologies. The main process is selectively fusing or melting the associated metal or alloy powder bed. To be specific, the effective layers of the particles are first completely melted by highenergy electron beams. Then, succeeding layers are melted to the preceding layer, forming repeated solid/liquid layered zones. It should be noted that this process is quite different from the traditional solidification process where the continuous melt occurs

Ti-6Al-4V lattice structures fabricated by electron beam melting for biomedical applications

1

2

3

4

5

6

7

8

9

10

279

11

Fig. 3.5.1 Eleven kinds of unit cells: (1) Square pyramid (G7); (2) Triangular prism; (3) Cubic; (4) Hexagonal prism; (5) Octagonal prism; (6) Rhombic dodecahedron; (7) Cuboctahedron; (8) Rhombicuboctehedron; (9) Truncated cube; (10) Truncated octahedron; (11) Truncated cuboctahedron.

at the solid/liquid (melting) interface. The corresponding processes are summarized in a group of diagrams as shown in Fig. 3.5.2A [21,22]. Nowadays, most EBM systems are from Arcam AB in Sweden, and Sciaky in the United States. Commercial application of this technology began in the 2000s. The EBM and welding system, in general terms, is similar to classical electronic scanning and transmission by electron microscopy configurations, except for a higher beam Rake Metal powder Start plate

Component Building tank

Filament Grid cup Anod

Electron beam

Process platform 1. Application of the powder layer

Focus coil Deflection coil Electron beam Powder container Vacuum chamber

4. Lowering of the process platform

2. Preheating of the powder layer

Building table

(B)

(A)

3. Melting of the cross section

Fig. 3.5.2 Schematic diagram of the process-fabricated Ti-6Al-4V scaffold by electron beam melting (A) and the internal components of electron gun and melting chamber (B) [22,23].

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current and a simpler lens system. A schematic description of an EBM system is shown in Fig. 3.5.2B. A high-energy electron beam (with the power of 4 kW) is generated in a standard electron gun configuration with accelerating potential operated at 60kV, corresponding to the beam current reaching high to tens of mill amperes. Unlike the selective laser melting (SLM) process (another widely used AM technology), the electron beam continuously scans the powder bed in a vacuum chamber, accompanied by the conversion of kinetic energy into internal energy. For the scan speeds, the EBM system is much greater than laser melting systems by an order of magnitude. Before each layer melting, a high temperature up to 700°C is needed to preheat the substrate plate by electron beam bombarding so as to reduce residual stresses and, more importantly, to sinter the powder to avoid powder smoking [23,24]. In addition, both the densification rate and microstructural homogeneity of EBM as-fabricated parts can be improved by optimizing technical parameters, which is extremely important to the mechanical properties of those as-fabricated parts. Many studies have been carried out on the performance of the EBM manufactured parts and improvement of the mechanical properties of the samples. By using an EBM system, a large number of implants have been manufactured; successful examples include knees, hip joints, jaws, and maxillofacial plate replacements [25–27]. The results showed that the EBM components provided the space for osseous tissue in-growth [28].

3.5.4

Surface characteristics and microstructure of Ti-6Al-4V cellular structures [23]

Fig. 3.5.3 shows the practical application of EBM technology in the fabrication of stochastic foam and reticulated mesh. Due to the fast cooling rate of the thin and isolated struts (Fig. 3.5.4A) [10,29], both ligaments and struts primarily consist of α0 -martensite. A small amount of smallish β phase can be detected by TEM, which was surrounded by thin lath acicular α0 -martensite [30] (Fig. 3.5.4C–F). The volume fraction of the β phase is too low for XRD to detect (Fig. 3.5.4B). The ligaments and struts have rather rough surfaces (Fig. 3.5.5).

3.5.5

Mechanical properties of Ti-6Al-4V cellular structures

3.5.5.1 The influence of porosity 3.5.5.1.1 Young’s modulus [30] There are two approaches to get Young’s modulus: one by the resonant frequencydamping analysis technique and the other by a compressive test. In order to match the modulus of surrounding bone tissues, which were 0.1–20 GPa, the effective way is to adjust the porosities of the Ti-6Al-4V foams and meshes from 50% to 95% [10,17,31] (Fig. 3.5.6). This can improve the mechanical properties by avoiding

Ti-6Al-4V lattice structures fabricated by electron beam melting for biomedical applications

281

(A)

(B) Fig. 3.5.3 Macroscopic images of the stochastic foams (A) and reticulated meshes (B) [31].

the stress-shielding and bone resorption. From the results, the modulus measured by resonant frequency damping is a little bigger than the ones tested by compress testing (Fig. 3.5.6A). The modulus of cellular structures and porosities follows a clearly linear relationship. Adjusting porosity is an effective way to get the desired modulus of foam and mesh. Gibson and Ashby successfully used the formulas to explain the relation of mechanical properties and the structures of porous materials [32]. The formula is used to describe the relationship of the relative modulus (E/Es) and the relative density (ρ/ρs) as:

Titanium in Medical and Dental Applications

50 μm

(A)

30

40

(B)

50 60 70 2q, degree

80

(202)

(103) (200) (112) (201) (004)

(110)

(102)

(100) (002)

Intensity

(101)

282

90

β

200 nm

(C)

(D)

200 nm

α⬘

(E)

200 nm

(F)

Fig. 3.5.4 Optical microstructures (A), XRD profile (B), and TEM images (C–F) of Ti-6Al-4V mesh struts [31].

E=Es ¼ ðρ=ρs Þ2

(3.5.1)

where the subscript “s” represents the parent materials of structure struts. Compared with classical models, the fitting exponential factor n, which is calculated from experimental results, is 2.0–2.4 and a little higher than the theoretical value (2.0) (Fig. 3.5.6B). Summed up in other reported results [13,15,16,19] and by our previous research [20] on stochastic foams, the fitting exponential factors are about 2.4–3.0. The main cause of this discrepancy is due to the different test methods chosen [10,33].

Ti-6Al-4V lattice structures fabricated by electron beam melting for biomedical applications

283

Fig. 3.5.5 SEM images of the Ti-6Al-4V foam ligament (A) and mesh strut (B) [31].

500 mm

(A)

500 mm

(B)

3.5.5.1.2 The compressive strength [17] According to the reported results, the compressive strength of Ti-6Al-4V structures with 50% porosity can reach about 300 MPa. The structures with 95% porosities are just 3 MPa (Fig. 3.5.7) [9,15,16,21]. During deformation, a straight crush band formed across the sample along an identical angle with the loading direction, where the compressive stress is max. However, for stochastic foams the crush band is tortuous and random [17,30]. There is a linear relationship between the specific modulus (E/ρ) and specific strength (σ/ρ); the specific strength (σ/ρ) decreases monotonically with decreasing specific modulus from the porosity of 62%–86% (Fig. 3.5.8). Compared to the stochastic foam, the regular mesh has higher specific strength. Ti-6Al-4V structures have better combinations of lower specific modulus and proper specific strength than other cellular structures [29,34] (Fig. 3.5.8).

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Young's modulus, GPa

102

Ti-6Al-4V feom Ti-6Al-4V mesh

E/Es

101

0.1 Cubic G7 Rhombic dodecahedron Rhombic dodecahedron

0.01

0

10

E/E0=(r/r0)n

10–1 0.0

1E-3 0.5

(A)

1.0 1.5 Density, g/cm3

2.0

0.1

1 r/rs

(B)

103

1

Cubic G7 Rhombic dodecahedron

G7 Rhombic dodecahedron

102 sp /ss

Compressive strength, MPa

Fig. 3.5.6 Young’s modulus of the meshes with different cell shapes (A), including dynamic Young’s modulus (open symbols) and static Young’s modulus (solid symbols). (B) Plots of the relative modulus (E/Es) versus relative density (ρ/ρs) for the reported Ti-6Al-4V foams and meshes [31].

0.1

1

10

100 0.4

(A)

0.8

1.2 1.6 Density, g/cm3

2.0

0.01 0.1

(B)

0.2 r /r s

0.3

0.4 0.5

Fig. 3.5.7 Compressive strength of the meshes with different cell shapes (A) and plots of relative strength (σ p/σ s) versus relative density (ρ/ρs) for the reported Ti-6Al-4V meshes (B) [18].

3.5.5.1.3 Compressive fatigue properties [23] The fatigue failure of Ti-6Al-4V scaffolds manufactured with the EBM technique is mainly caused by gradual strain accumulations and the abrupt eventual collapse [23]. In order to ensure the safety of long-term implants, the fatigue strength should be considered. But the fatigue strength of Ti-6Al-4V scaffolds that have ratios ranging from 0.1 to 0.2 is still a little lower than the parent materials, which is about 0.6. This owes to the rough surface and brittle α’-martensite (Fig. 3.5.9A) [34]. Like reported nickel and aluminum foams, the Ti-6Al-4V scaffolds show a better linear relationship (Fig. 3.5.9B).

E/r, GPa/(g/cm3)

Ti-6Al-4V lattice structures fabricated by electron beam melting for biomedical applications

10

Fig. 3.5.8 Relation between the specific modulus (E/ρ) and the specific strength (σ/ρ) of the cellular Ti-6Al-4V alloy with foam and mesh structures, and in comparison with other cellular metallic materials summarized by Ashby et al. [31].

Cymat Al-Si Alulight Al Alporas Al ERG Al Inco Ni Foam Ti-6Al-4V Mesh Ti-6Al-4V

1

0.1 0.1

285

1

10

100

100

(A)

100 Fatigue strength, MPa

Compressive strength, MPa

s /r, MPa/(g/cm3)

10

1

0.1 104

0.73 g/cm3 0.91 g/cm3 1.12 g/cm3 1.68 g/cm3

Ti-6Al-4V mesh Al foam Al-SiC foam Ni foam

10 1 0.1 0.01

105

106

107

Number of cycles to failure

0.01

(B)

0.1

1

10

100

Young's modulus, GPa

Fig. 3.5.9 S-N curves of the Ti-6Al-4V mesh arrays with different density (A) and plots of compressive fatigue strength versus Young’s modulus (B) [24].

3.5.5.2 The influence of cell shape Three kinds of cellular structures were considered: cubic, G7, and a rhombic dodecahedron (Fig. 3.5.10). The strut thicknesses are about 700 μm, which is 200 μm thicker than the design value (Fig. 3.5.10G–I).

3.5.5.2.1 Young’s modulus [17] A dynamic Young’s modulus was presented in Fig. 3.5.4A. According to the results, the modulus of these scaffolds increases from 0.5 GPa to 15 GPa with density increases. Through the choice of proper porosities, these scaffolds are comparable to trabecular, which is about 0.05–3 GPa, and cortical bone tissues, which are about 10–25 GPa [35,36]. Fig. 3.5.6A also shows that the modulus of cubic scaffolds is the highest while G7 is the lowest.

286

Titanium in Medical and Dental Applications

(A)

(B)

(C)

(D)

(E)

(F)

(G)

(H)

(I)

Fig. 3.5.10 Cubic, G7, and rhombic dodecahedron element in the Materialise software (A–C), the corresponding Ti-6Al-4V prototype blocks fabricated by EBM method (D–F) and SEM images of the meshes (G–I) [18].

3.5.5.2.2 Static compressive properties [17] The compressive stress-strain curves of the three structures all consist of three parts: elastic deformation until to the compressive strength, a plateau region, and a densification region. In this part, the stress increases quickly with the strain. As is well

Ti-6Al-4V lattice structures fabricated by electron beam melting for biomedical applications

287

known, the structures with stress-strain curves have obvious fluctuation characteristics in plateau regions that are referred to as brittle failure modes [32], just like the cubic and rhombic dodecahedron structures (Fig. 3.5.11A and C). However, the G7 structure shows a ductile failure mode where stress-strain curves have a smooth plateau region (Fig. 3.5.11B). These results illustrate that through optimization of cellular structures, they exhibit different deformation behaviors. In situ SEM observations were used to detect the deformation of struts in different meshes (Fig. 3.5.12). The strut deformation of a cubic structure is buckling. For cubic structures, the deformations of struts are buckling, and for G7 and rhombic dodecahedron structures, the deformation behaviors are coupling of buckling and bending. The crack usually initiates in the connections of struts. The mechanical properties are summarized as follows. The compressive strength of cubic, G7, and rhombic dodecahedron structures ranges from 10 to 300 MPa, as shown in Fig. 3.5.7A. Furthermore, the compressive strength of cubic scaffolds is the highest while G7 is the lowest with the same densities. According to the porous materials Gibson-Ashby model [32]: E=Es ¼ ðρ=ρs Þ2

250

1.6 g/cm3 0.9 g/cm3 0.7 g/cm3 0.6 g/cm3 0.5 g/cm3

200

200 Stress, MPa

Stress, MPa

300

(3.5.2)

100

1.8 g/cm3 1.6 g/cm3 1.0 g/cm3 0.8 g/cm3

150 100 50

0 0.0

(A)

0.2

0.4

0.6

0.8

0.2

(B)

Strain, mm/mm 140

0.4

0.6

0.8

Strain, mm/mm

0.62 g/cm3 0.73 g/cm3 0.91 g/cm3 1.18 g/cm3 1.68 g/cm3

120 Stress, MPa

0 0.0

1.0

100 80 60 40 20 0 0.0

(C)

0.2

0.4

0.6

0.8

1.0

Strain, mm/mm

Fig. 3.5.11 Nominal compressive stress-strain curves of the meshes with cubic (A), G7 (B), and rhombic dodecahedron (C) structures [18].

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P

500 mm

(A)

P

A

A

A

B

B

B

e = 0%

e = 0.5%

P

P

D 500 mm

(B)

e = 1.5%

P

D C

P

e = 0%

D C

C e = 1%

e = 2%

Fig. 3.5.12 In situ SEM observations of the meshes with cubic (A) and G7 (B) cells at different strains during the compression [18].

σ p =σs ¼ Cðρ=ρs Þ1:5

(3.5.3)

σ p =Es ¼ C0 ðρ=ρs Þ2

(3.5.4)

Where the subscript “s” represents the parent materials of the structure struts. C and C0 are the constant values determined by strut materials. This mode can be used with other kinds of metallic porous materials [32]. However, there is a lack of actual regular structures to calculate these parameters. The Gibson-Ashby models for three kinds of scaffolds are shown as follows: For the cubic cell, E=Es ¼ C1 ðρ=ρs Þ

(3.5.5)

σ p =Es ¼ C2 ðρ=ρs Þ2

(3.5.6)

For the G7 cell, E=Es ¼ C3 ðρ=ρs Þ2

(3.5.7)

σ p =σ s ¼ C4 ðρ=ρs Þ1:5

(3.5.8)

Ti-6Al-4V lattice structures fabricated by electron beam melting for biomedical applications

289

For the rhombic dodecahedron cell, E=Es ¼ C5 ðρ=ρs Þ2

(3.5.9)

σ p =σ s ¼ C6 ðρ=ρs Þ1:5

(3.5.10)

According to the relative modulus and densities of three kinds of scaffolds, the fitting exponent n of the cubic structure is 2.4, for the G7 structure is 2.0, and for rhombic dodecahedron structures is 2.2, respectively (Fig. 3.5.13A). The fitting exponents of G7 and rhombic dodecahedron structures are closest to the theoretical value; their struts are dominated by the bending deformation. For the cubic structure, the deformation of their struts is buckling; however, the obvious discrepancy between the experimental factor (2.4) and the theoretical one (1) may be caused by the uniform struts. For the relative strength and densities of three kinds of scaffolds, the fitting exponential n0 of the cubic structure is 1.7, for the G7 structure is 1.9, and for the rhombic dodecahedron structure is 2.2, respectively (Fig. 3.5.7B, Fig. 3.5.13C). There exists a difference between fitting factors and theoretical values. Such discrepancies can be explained by the following reasons. On one hand, the brittle martensitic phase and rough surfaces contribute to the brittle failure. On the other hand, on the G7 and rhombic dodecahedron structures, just the bending component of the struts was considered in the theoretical analysis. Actually, the deformation behavior of the struts is the coupling effect of the bending and buckling components, especially for rhombic dodecahedron structures.

3.5.5.2.3 Compressive fatigue properties [37] According to early analysis, the structure of scaffolds highly determines the static mechanical properties. The mechanical properties and deformation behaviors are significantly dependent on the structure parameters, for example, porosities, pore sizes, and cellular structures [17]. There are some results that illustrate the influence of 10

sp /Es

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0.1 0.2 r/rs

0.3

0.4 0.5

0.1

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0.2

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0.4

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r/rs

Fig. 3.5.13 Plots of the relative modulus (E/Es) (A), relative strength (σ p/σ s) (B), and versus relative density (ρ/ρs) for the studied Ti-6Al-4V meshes [18].

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surface morphologies, defect distributions, and loading conditions on the fatigue strength of scaffolds [38]. However, their underlying fatigue mechanism has not been clarified.

3.5.5.3 Fatigue mechanism 3.5.5.3.1 Cyclic ratcheting

10

10

1

1

0.1 0.01

G7-cyclic ratcheting G7-fatigue damage Cubic-cyclic ratcheting Cubic-fatigue damage RD-cyclic ratcheting RD-fatigue damage

1E-3

Strain, %

Strain, %

Under compressive loading, the struts deformed by buckling and/or bending, which can result in the strain accumulation of the whole sample [39]. In Fig. 3.5.14, three kinds of scaffolds have taken the fatigue strain accumulation during the fatigue progress, especially high cycle fatigue. Under identical stresses, the cubic structure has the smallest cyclic ratcheting rate (dε/dN) and the G7 structure has the biggest (Fig. 3.5.14C). According to the aforementioned static compression results, the fatigue behavior of the studied meshes was discussed. First, when the strut deformed mostly by buckling, the stress on the strut was the compressive loading. Compared to the tensile stress, the

0.1 0.01

G7-cyclic ratcheting G7-fatigue damage Cubic-cyclic ratcheting Cubic-fatigue damage RD-cyclic ratcheting RD-fatigue damage

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40

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100 120 ⫻1000

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20

40

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1E-5 1E-6 1E-7 Cubic G7 (Materialise/Magics) Rhombic dodecahedron

1E-8 1E-9 10

(C)

100 Stress, MPa

1000

Fig. 3.5.14 The cyclic ratcheting and fatigue damage strains of the meshes with cubic, G7, and rhombic dodecahedron cells in low (A) and high (B) fatigue cycle regions. RD in the figure is the abbreviation for rhombic dodecahedron. (C) Effects of stress on the cyclic ratcheting rate of the meshes with different structures [38].

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compressive stress promotes the closure of fatigue cracks [40]. Second, meshes dominated by buckling exhibited higher compressive strength. Compared to the meshes dominated by bending, they have the smallest deformation strain at identical stress levels; therefore it is easy to deduce that they have the lowest cyclic ratcheting and thus the highest fatigue endurance.

3.5.5.3.2 Fatigue damage of the struts During fatigue, the complex structures easily lead to stress concentrations and cause the initiation of the fatigue cracks. When the struts were under the coupling deformations of buckling and bending, according to Fig. 3.5.12B, the fatigue damage became the dominant factor influencing the fatigue behavior of the Ti-6Al-4V scaffolds (Fig. 3.5.14A and B). This influence on fatigue behaviors was more significant during the low cycle fatigue progress (Fig. 3.5.14A). In this case, the factors influencing fatigue crack initiation and propagation, such as surface properties, pore defects, and stress distribution in the struts, became more important in determining the fatigue properties of the cellular structure, which deteriorated the fatigue resistance [39,41]. According to the above results, even though the surfaces of cubic structures were rougher than G7 and rhombic dodecahedron structures, the fatigue strength of the cubic structure was still higher than them. It demonstrated that structures were the dominant factor on scaffold fatigue. Heat treatment is another way to improve the fatigue property of scaffolds by ameliorating the mechanical properties of struts [23,42].

3.5.6

Biocompatibility of Ti-6Al-4V cellular structures [43]

A 3D cellular implant is supposed to be a good alternative to bone regeneration and hard tissue due to its advantage in providing the necessary support for cell proliferation while maintaining its differentiation function and replacement [44,45]. By controlling their pore size, pore geometry shape, and porosity distribution, the cellular structures can significantly adjust the cells in in vitro and in vivo environments, and good biocompatibility can be achieved [46–49]. For instance, Ponader et al. [50] evaluated the applicability of different Ti-6Al-4V surfaces produced by electron beam smelting processes as an accessory matrix for human fetal osteoblast proliferation and differentiation. It was found that osteoblast proliferation and differentiation was influenced by the surface characteristics of Ti-6Al-4V meshes, which can be controlled by adjusting the process parameters. By using chemical surface modification, the surface biological activity of EBM-fabricated porous Ti and Ti-6Al-4V components is improved [51], leading to an easier fixation around the implant bone and improved long-term stability [52,53]. Meanwhile, the in vivo test on sheep anterior cervical fusion efficacy manifested a fast bone in-growth and a better osseointegration as well as a superior mechanical stability for the porous Ti cage fabricated by EBM, compared with the conventional polyetheretherketone (PEEK) cage. Thus, a 3D cellular implant is believed to have great potential for clinical

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(A)

(B)

(D)

(E)

(F)

Ti-6Al-4V

6 months

6 months

(C)

Fig. 3.5.15 Porous EBM Ti-6Al-4V cage (A) and PEEK cage (B) used in animal tests. Histological images of PEEK cage and Porous EBM cage over postsurgery recovery time. (C) and (D) are Porous EBM cage (6 months), bone matrices in the pores were found mainly around the struts and lost their normal cancellous appearance. Intimate bonding between metal and bone matrix was observed. (E) and (F) are PEEK cages (6 months), the gap was still obvious and an intimately bonded interface between the bone matrix and the material was rarely seen [55].

application (Fig. 3.5.15) [54]. With the same technology, Li et al. [55] prepared porous titanium alloy rods and implanted them into a sheep, recently. The in vivo tests indicated that the mechanical properties and biocompatibility of porous titanium alloy rods made by EBM met all the requirements when used in the early stages of treatment.

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Open-cellular structural implants, including mesh and foam, show many evidencebased advantages in addition to being able to tailor the structure to reduce the stress shielding effect in bone-compatible implantation. First, implant fixation enhances the production of bone cells (osteoblasts), which on one hand eliminates the need for cement and on the other hand ensures bone integration. Second, open-cellular structures can help eliminate or control the infection effect with their convenience in inserting antibiotics [56]. Third, quite different from the conventional machined implants illustrated for the entire knee arthroscopy shown in Fig. 3.5.16, porous implants fabricated by EBM are individualized, which means implants can be designed and prepared for a specific patient from micro computed tomography (micro-CT) scans. Illustrated examples are displayed in Figs. 3.5.12–3.5.14. More importantly, EBM-fabricated porous implants can be functionally graded to match soft core-hard shell bone structure where the interior soft bone stiffness is low to 2 GPa and may also be conducive to vascularization while the outer hard bone stiffness is high to 20 GPa [57]. According to Karageorgiou and Kaplan [31], the pore size of open-cellular implants is suggested to be greater than 300 μm, which can help to support bone formation and allow for potential vascularization. Recently, Nune

Fig. 3.5.16 X-ray images of a total knee replacement using commercial implants in a knee, f and t indicate femoral and tibial implants, respectively [44].

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et al. [58] suggested that pore sizes ranging from 400 to 800 μm (Fig. 3.5.17) are favored for cell differentiation occurring at an early stage while contributing to bone cells adhering to the struts and ligaments by forming a mass of cytoplasmic extensions. Within about 21 days, bone cells migrated through the cellular structure, finally bridging the pores [58]. These Ti-6Al-4V cellular structures, with the corresponding stiffness ranging from 2 GPa to 20 GPa, cover the functional stiffness range of bone structures, preventing the stress shielding phenomenon illustrated in Figs. 3.5.18 and 3.5.19.

Fig. 3.5.17 Tibial (knee) stem designs in (A) and (B) compared with an EBM-fabricated prototype shown in (C). (A) and (B) are CAD models rotated 45 degrees relative to one another [44].

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(A)

(B) Fig. 3.5.18 CAD models for open-cellular foam structures showing functionally graded porosity. (A) shows overall perspective while (B) shows top view [44].

3.5.7

Conclusion and future research trends [43]

Several living examples of Ti-6Al-4V implants fabricated by the EBM layer manufacturing process were given in Fig. 3.5.20. Although rich results have been achieved on mechanical properties and biocompatibility of evidence-based medicine with Ti-6Al-4V cell products, there is still some work to be done. Some of this is introduced below: (1) Cellular structures with graded/gradient porosity.

Until now, most studies mainly focused on open-cell structures and foams prepared with traditional metal titanium with a uniform pore size and a uniform relative density. However, in order to promote bone growth, the higher the porosity for the cell structures the better, which will unfortunately worsen their mechanical properties [59].

296

Fig. 3.5.19 (A) Simulated rod inserted into femur and (B) fabricated (EBM) component with cut away section to show functional porosity corresponding to Fig. 3.5.17A [44].

Fig. 3.5.20 Examples of experimental biomedical replacements produced by EBM layer manufacturing.

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Meanwhile, the uniform porous materials, and thus the unchanged modulus and stress concentration, will inevitably cause periprosthetic early failure of the implant [60]. Considering these shortcomings, artificial cellular Ti and Ti alloys with density or structure gradients (Figs. 3.5.18 and 3.5.19) are being developed. By adjusting structural parameters, including the distribution of pore size, relative density, shape, etc., the open-cellular structures are supposed to possess both high porosity and high strength, and may match the multiple functions of natural bone tissues. Although some attempts have been made to fabricate porous Ti alloys, some topics—such as biomedical applications [59], possibilities, design rules, evidence-based medicine manufacturing techniques, and the underlying mechanisms during loading—are still worth investigating. At the same time, structural graded porous titanium alloys still need further studying. (2) New β-type cellular titanium alloys comprising nontoxic and nonallergic elements.

Ti-6Al-4V has been widely used for orthopaedic implants. However, the toxicity of alloying elements Al and V has long been noted [61]. Thus, low modulus β-type titanium alloys are currently being developed by using nontoxic, nonallergenic elements [2,61] such as Ti-13Nb-13Zr, Ti-29Nb-13Ta-4.6Zr, Ti-35Nb-5Ta-7Zr, Ti-24Nb-4Zr8Sn (Ti2448), etc. Most of them mainly contain a large amount of Nb, Ta, and Zr. Owing to attractive potential applications of the AM techniques in biomedical fields, the processing-microstructure-property relationship of these new β-type cellular titanium alloys fabricated using the EBM technique should be further evaluated [47], especially because the modulus value of these alloys in their pore-free form can be less than 50 GPa compared with about 110 GPa for Ti-6V-4V.

Acknowledgments This study was supported partially by Chinese MoST Projects (2015AA033702, 2016YFC1102601).

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Additive manufacturing of cp-Ti, Ti-6Al-4V and Ti2448

3.6

T.B. Sercombe*, L.-C. Zhang†, S. Li‡, Y. Hao‡ *University of Western Australia, Crawley, Australia, †Edith Cowan University, Perth, Australia, ‡Shenyang National Laboratory for Materials Science, Institute of Metal Research, Chinese Academy of Sciences, Shenyang, China

3.6.1

Introduction

Despite titanium being the ninth-most common element in the earth’s crust and the fourth-most abundant structural metal (after iron, aluminum, and magnesium), the high cost of reducing and refining the ore makes titanium an expensive material. Fortunately, it does possess an attractive set of properties, including high strength, low weight, and excellent corrosion resistance. Nonetheless, titanium is only used in applications where the high cost can be justified or at least tolerated. For example, the high specific strength makes the metal attractive to the aerospace industry while the high corrosion resistance sees it used in chemical processing equipment. In its commercially pure form, and in some titanium-based alloys, it is also biologically compatible in humans, which, combined with the corrosion resistance and high strength, make it attractive for biomedical devices. A summary of the characteristics of elemental titanium is shown in Table 3.6.1. At temperatures below 882°C, titanium exists in a hexagonal close-packed (HCP) phase, known as alpha (α) phase. Above 882°C, this transforms to a body-centered cubic beta (β) phase. Different alloying elements affect the stability of the α and β phases. Classified as α-stabilizers, β-stabilizers, or neutral elements, they, along with the cooling rate, play a crucial role in determining the room temperature microstructure. Commonly used α-stabilizers include aluminum and the interstitial elements oxygen, nitrogen, and carbon; this results in the retention of α at temperatures >882°C. In contrast, β-stabilizers such as Nb, Mo, and V result in the stabilization of the beta phase at low temperature. If added in sufficient quantity, the β phase can be retained to room temperature. As their name suggests, neutral elements have no effect on the stability of the phases.

3.6.1.1 Titanium as an implant material The ideal load-bearing implant should provide a long service life, have sound mechanical properties (such as high strength, wear, and fatigue resistance), high corrosion resistance, and an elastic modulus closely matched to that of bone (the structure to be

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Table 3.6.1

Titanium in Medical and Dental Applications

Properties of elemental titanium Titanium properties

Atomic number Atomic weight Crystal structure (α) Crystal structure (β) Density Melting point Boiling point Specific heat (at 25°C) Heat of fusion Poisson’s ratio Coefficient of thermal expansion

22 47.9 g/mol HCP (T < 882.5°C) BCC (T > 882.5°C) 4.51 g/cm3 1668°C 3260°C 0.52 J/kg K
440 kJ/kg 0.41 8.64  106/°C

From M. Donachie, Titanium A Technical Guide, ASM International, OH, 1988.

strengthened or replaced) [1,2]. The large difference between the modulus of elasticity of the implant and that of the bone causes stress shielding, which leads to loosening of implants and painful revision surgery [3]. Most importantly, the implant materials used must not exhibit cytotoxicity to the recipient [4]. This last constraint places limitations on the alloying elements used for the purpose of implants in humans. With advancements in medicine and the associated increase in life expectancy, the life span of an implant is increasing [3]. To avoid the costly and painful requirement for additional surgeries or supplementary repair, the longevity of the implant is of high importance [2]. With the instances of hip revision surgeries expected to increase by 137% between 2005 and 2030 and knee revisions by 602% in the same period [2], it is important that biomaterial development be a priority. While materials such as 316L stainless steel (SS) and cobalt-based alloys are commonly used as biomaterials, titanium and its alloys are emerging as the first choice for many applications as they are biocompatible, highly resistive to corrosion, and have high strength-to-weight ratios. The modulus of SS and chromium-cobalt alloys (210 and 240 GPa, respectively) are also much greater than that of bone, leading to stress shielding, while titanium alloys have lower moduli, ranging between 112 and 55 GPa. Table 3.6.2 lists some common biomedical alloys and their moduli.

3.6.1.2 Commonly used titanium alloys The relatively low modulus of titanium, along with its excellent corrosion resistance and very high specific strength, has resulted in it being widely favored for bonereplacement applications [5]. Of the vast array of titanium alloys available, only a limited number are used for orthopedic applications. The ubiquitous Ti-6Al-4V is the most widely used. However, the cytotoxicity of the vanadium, both in the elemental

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Young’s Modulus of various of biomedical alloys compared to that of bone

Table 3.6.2 Alloy

Young’s modulus (GPa)

CoCr (cast) ANSI 316L stainless steel Tantitium Ti-6Al-4V Ti-6Al-7Nb Commercially pure Ti Beta-titanium alloys Cortical bone

240 210 200 112 110 100 40–60 10–30

Adapted from R. Yang, S. Li, Y. Hao, Development and Application of Low-Modulus Biomedical Titanium Alloy Ti2448, Intech Open Access Publisher, 2011.

state and in the form of an oxide, as well as the adverse effect of vanadium and aluminum ions on the body continue to be problematic [6,7]. As a consequence, an alternate alloy, Ti-6Al-7Nb, has been developed to specifically overcome the potential of vanadium toxicity. The superior corrosion resistance of commercially pure titanium (or cp-Ti) means that it is also relatively widespread, especially where the loads are relatively low (e.g., cranial plates). The one characteristic that these alloys share is a relatively high modulus compared to that of bone. Although the modulus of cp-Ti, Ti-6Al-4V, and Ti-6Al-7Nb is lower than that of other biomaterials (see Table 3.6.2), it is still significantly above that of bone. A mismatch of moduli between the biomaterial and surrounding bone can cause stress shielding in the bone, which eventually leads to bone resorption. This has been identified as a major causal factor of implant loosening [8,9]. The skeleton in our body is under load, both from body weight and the action of our muscles. As the loading conditions change from site to site, so too does the density and therefore the strength of the bone. Where the loads are high (e.g., femur, mandible) the bone is at its densest. Where there are only low loads, (e.g., fingers), the density is much lower [10,11]. Indeed, in healthy people, an increase or decrease in activity can also affect the density of their bones. Known as Wolff’s Law [12], the same changes can occur as a result of stress shielding, which can be caused by the insertion of a high stiffness implant. The high stiffness metal implant takes the majority of the load with a corresponding drop in the load going through the surrounding bone. Such changes in load sharing can cause remodeling of the bone to a lower density, which can significantly affect the longevity of the device. One way to reduce the chances of stress shielding is to lower the modulus of implant materials. As such, a range of low modulus-beta titanium alloys has been developed that has a modulus approximately half that of conventional titanium. One of the most promising of these alloys is Ti2448, a titanium alloy containing by weight 24% Nb, 4% Zr, and 8% Sn with the balance Ti.

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3.6.2

Titanium in Medical and Dental Applications

Additive manufacturing

3.6.2.1 Overview Additive manufacturing (AM), historically referred to as rapid prototyping, is a range of manufacturing technologies characterized by depositing material in successive layers to form a three-dimensional (3D) object. The AM process begins with a computer-aided design (CAD) model of the desired part, which is then sliced and fed into the AM machine in order to direct the machine to print material in the correct area [13]. Conventional and traditional methods of manufacturing often involve fabrication of parts by removing material from a larger stock or sheet metal; however, AM only deposits material where required, reducing waste. Furthermore, there are some complex geometries that are costly and/or impossible to produce with conventional or subtractive manufacturing methods. The layer-wise, additive nature of AM allows for complex geometries to be produced without considering design for manufacturing and assembly, ultimately encouraging design innovation [14]. Since the first AM patent by Chuck Hull for a technique he called stereolithography (SLA) in 1986 [15], there have been many developments. Currently, there are several main AM technologies and these are summarized in Fig. 3.6.1. Of these, selective laser melting (SLM) and electron beam melting (EBM) are the most promising for orthopedic implants.

3.6.2.2 Selective laser melting SLM is a method of AM that uses a high-power laser beam to selectively melt regions of a metal powder bed. Successive layers of melted metal powder create a 3D-object. When compared with selective laser sintering (SLS), the SLM process melts the powder feedstock completely while SLS does not. The complete melting of the powder allows for fusion to occur between the layers of the build [16], resulting in high density and good mechanical properties.

3.6.2.2.1 The SLM process Fig. 3.6.2 illustrates schematically the SLM process. The process begins with a CAD model of the desired part, which is then sliced into printable layers using computer software. The physical production of the model begins with the deposition of a thin layer of powder onto a substrate within a build chamber filled with inert gas, usually Ar or N2. The inert gas reduces oxidation during the build process. Using a pattern defined by the CAD model, the powder is then melted with a high power fiber laser (λ ¼ 1.06 μm). The melted powder rapidly cools into a solid layer and the platform is then lowered. A new layer of powder is applied and the process is repeated. When the process is complete, the solid object is left among the powder and can be removed for further processing as required.

Liquids

Metals

Ceramics

Molten materials

Photopolymers Cure via vector scanning (laser)

Cure via raster scanning (projection)

Cure via lamp after selective deposition

Extrusion

Ink Jetting

SLA (Stereolith ograpgy)

DLP (digital light projection)

MJM (multi-jet modelling)

FDM (fused deposition modelling)

DoD (drop on demand)

Powders

Polymers

Coaxial laser deposition

Powder bed fusion Fused by e-beam

EBM (electron beam melting)

Fused by laser

SLS (selective laser sintering)

SLM (selective laser melting)

Fused via binder

BJ (binder jetting)

Additive manufacturing of cp-Ti, Ti-6Al-4V and Ti2448

Additive manufacturing

LMD (laser metal deposition)

Fig. 3.6.1 Summary of Additive Manufacturing Technologies. Adapted from N. Kang, et al., Microstructure and strength analysis of eutectic Al-Si alloy in situ manufactured using selective laser melting from elemental powder mixture, J. Alloys Compd. 691 (2017) 316–322.

307

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Laser source Scanner mirrors X-Y deflection Powder scraper

f-q lens Melt pool

Y

X Feed container Base plate Build platform Overflow container

Z

Fig. 3.6.2 Schematic of the SLM process [17].

The energy applied by the laser is a characteristic that is typically measured and used to compare different build strategies. The energy supplied to a volumetric unit of powder is defined by: E¼

P vst

(3.6.1)

where P is the laser power in watts, v is the last scan speed in mm/s, t is the layer thickness in mm, and s is the scan spacing in mm. Other factors such as scanning strategy and powder characteristics are also important factors in the success of a build.

3.6.2.2.2 SLM issues The SLM process often leads to the formation of defects including cracks, delamination, porosity, and poor surface roughness. Although the surface roughness can sometimes be improved through the use of expensive postprocessing, this can be problematic for highly complex parts. Porosity is a common defect among metal AM parts and can adversely affect the mechanical properties of the part, especially the fatigue life. Porosity can be powder induced or process induced. Feedstock powder may contain gas pores as a result of the production method. These gas pores can translate directly to the SLM built part. Process-induced porosity can be a result of the applied energy not being sufficient enough to completely melt the powder, or being too high, resulting in vaporization. Through optimizing the scanning strategy, it is possible to reduce the porosity to a very low level [16]. Residual stress is the stress within a material that remains present after the removal of any applied loads. When the residual stress exceeds the yield strength of a material, deformation may occur. Due to the localized and rapid temperature changes of the

Additive manufacturing of cp-Ti, Ti-6Al-4V and Ti2448

309

SLM process, temperature gradient-induced residual stress is a common problem in parts produced by SLM [16]. Although heating the powder bed may reduce the effects of residual, it is not always completely effective. Postbuild stress-relieving heat treatment (with the parts still attached to the substrate plate) is often undertaken in order to minimize the amount of distortion. Cracking and delamination are defects that can occur within the material as a result of the high cooling rate and/or residual stresses. Cracking can occur either during solidification or subsequent heating and is most problematic in materials with low inherent ductility. Cracking can also be initiated by pores, especially in the presence of significant residual stresses. Delamination occurs when subsequent layers separate due to insufficient bonding between layers. It can be caused by incomplete melting between the layers or insufficient remelting of the underlying solid.

3.6.2.3 Electron beam melting The fundamental difference between SLM and EBM is the source of the energy used to melt the powder. As shown schematically in Fig. 3.6.3, EBM uses a high-power electron gun to melt the powder. Another significant difference between the two technologies is that EBM is performed in a vacuum of about 104 to 102 mbar during the build phase. In addition, SLM uses galvanometer-controlled scanning mirrors to direct the laser beam while in EBM the electron beam is managed by electromagnetic coils that provide extremely fast and accurate beam control and allow several melt pools to be maintained simultaneously. EBM has several advantages over SLM. There is a high efficiency in converting the electrical energy into electron beam energy. In addition, almost all the electron energy is absorbed by the metal powder, resulting in efficient energy transfer, fast scanning speeds, and high penetration depth. However, the powder needs to be heated to a high temperature to lightly sinter the particles together in order to prevent powder from being blown away from the part bed [18]. To achieve this, the beam is scanned multiple times across the part bed using high speed (104 mm/s) and high current (30 mA) prior to the final melt scan [4]. Fortunately, this is easy to achieve with the high energy and rapid scanning ability of the electron beam. The final melt scan is performed at a slower speed (102 mm/s) and lower beam current (5–10 mA). As a result of the high preheat temperature, EBM parts experience a slower cooling rate and therefore a different microstructure than SLM. This is particularly true with Ti-6Al4V, with the martensitic α0 phase being favored in the higher cooling rate SLM with the more traditional dual phase α + β occurring in EBM [19]. Although a range of different materials has been successfully produced using EBM, including 316L [20], TiAl [4], Co-Cr-Mo alloys [21], and nickel-based superalloys [22], by far the most widely used is Ti-6Al-4V. The structure-processingproperty relationships have been widely studied on Ti-6Al-4V produced via EBM on both solid and porous lattice structures. Factors such as mechanical [23] and fatigue [24] properties, corrosion resistance [25], and surface roughness [26] have all been topics for investigation.

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Filament Cathode −60 kV Gate (0 ~ −1500 V to the cathode)

Electron beam gun

Anode

Focusing coil

Deflection coil

Vacuum chamber

Powder storage Powder distributor Powder bed

Electron beam Part Substrate Building platform

Fig. 3.6.3 Schematic representation of the EBM equipment. From C. Guo, W. Ge, F. Lin, Effects of scanning parameters on material deposition during electron beam selective melting of Ti-6Al-4V powder, J. Mater. Process. Technol. 217 (2015) 148–157.

3.6.3

Additive manufacturing of Ti2448

3.6.3.1 Selective laser melting of solid parts Unlike the widely studied AM of Ti-6Al-4V, very few studies have been conducted on low modulus beta titanium alloys due to the absence of easily obtained powder. Zhang et al. [27,28] have studied the densification, microstructure, and mechanical behavior of Ti-24Nb-4Zr-8Sn (Ti2448) manufactured by SLM. Both Vickers microhardness and relative density of SLM-produced Ti2448 samples are closely related to the laser processing parameters (Fig. 3.6.4). The relative density generally increases with the decrease in laser scan speed up to 600 mm/s and it tends to plateau at >99%. The critical laser energy density for SLM of near fully dense Ti2448 is about 33 J/mm3 [28], which is about one-third that of cp-Ti and Ti-6Al-4V (both of which are around 120 J/mm3 [29]). Near full density parts (>99%) have been obtained at a laser power of 200 W and a scan speed range of 300–600 mm/s. The microhardness reached

Additive manufacturing of cp-Ti, Ti-6Al-4V and Ti2448

311

240 220

95

200 90 180

Relative density Vickers hardness 85

80

160 laser power: 200 W 200

300

400

Vickers hardness (Hv)

Relative density (%)

100

140 500

600

700

800

900

Laser scan speed (mm/s)

Fig. 3.6.4 Relative density and Vickers hardness of the SLM-produced Ti-24Nb-4Zr-8Sn as functions of different laser scan speeds [27].

220 HV for the near fully dense parts. The tensile properties of the Ti2448 samples fabricated by SLM, hot rolling, and hot forging are compared in Table 3.6.3 [27]. The Young’s modulus and ductility of all samples are comparable. On the other hand, the yield and ultimate tensile strengths of SLM-produced samples are slightly lower than those produced by rolling and forging. This is most likely due to the fact that parts were loaded parallel to the build direction, which usually produces the lowest properties [27,32]. Furthermore, the SLM-produced Ti2448 samples do not show the pronounced superelastic behavior as the traditionally processed counterparts. This suggests that the nanoscale Nb modulation created by phase decomposition would be suppressed by the fast cooling [33–35].

Comparison of the tensile properties for the Ti-24Nb-4Zr-8Sn alloys manufactured by selective laser melting (laser powder: 200 W; laser scan speed: 550 mm/s) and by conventional processing methods [27]

Table 3.6.3

Processing methods

E (GPa)

σ 0.2 (MPa)

σ UTS (MPa)

δ (%)

Reference

Selective laser melting Hot rolling Hot forging

53  1

563  38

665  18

13.8  4.1

[27]

46 55

700 570

830 755

15.0 13.0

[30] [31]

Young’s modulus E, yield strength σ 0.2, ultimate tensile strength σ UTS, and elongation δ.

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3.6.3.2 Production of porous structures Although the Young’s modulus of the Ti2448 alloy (42–50GPa) is relatively low, it needs to be further reduced if it is to match that of bone (4–30GPa) [36,37]. One common method to reduce the modulus of a material is to introduce porosity into the structure [38–41]. There are two main reasons for this. First, introducing porosity will further decrease the modulus of titanium closer to that of human bone, thereby minimizing the stress shielding effect. Secondly, porous structures enhance biological fixation through bone cell in-growth between implant and adjacent bone and therefore facilitate long-term fixation of the implant [42]. The geometric freedom offered by AM technologies such as SLM and EBM is considered one of the most promising methods for fabricating the complicated porosity structure of artificial bone implants [43–45].

3.6.3.2.1 SLM of porous titanium structures SLM has been widely used to produce porous structures from many types of titanium alloy materials and composites [29,46–49]. Furthermore, the design and manufacture of novel titanium structures for improving bone in-growth has also been studied [41,50]. It has been demonstrated that SLM is able to produce optimized structures ideal for bone in-growth. SLM-fabricated porous cp-Ti structures with 55%–75% porosity (a level analogous to human cancellous bone) exhibit a compressive strength between 35 and 120 MPa [51]. Moreover, SLM-produced porous cp-Ti and Ti-TiB composite materials with porosity levels of 10%, 17%, and 37% show yield strength and elastic modulus in the range of 113–350 MPa and 13–68 GPa for cp-Ti and 234–767 MPa and 25–84 GPa for Ti-TiB composite materials, respectively [48]. Such low values are close to that of human bone, indicating that they can be considered as a potential candidate for biomedical implants. Extensive endeavors have also been made to study the processing and properties of SLM-produced porous Ti-6Al-4V alloys. For example, Warnke et al. [52] produced Ti-6Al-4V scaffolds using SLM for potential use in bone tissue engineering applications. They found that the biocompatibility to human osteoblasts was very good and that this was coupled with high compressive strength. This work also showed that SLM was able to reproduce complex microscopic features from the original designs. The average compressive modulus of tested samples was 2.97 GPa, which is between that of trabecular bone (0.1–0.5 GPa) and cortical bone (15 GPa). Furthermore, theoretical and experimental measurements were conducted on octahedral Ti-6Al-4V porous structures in terms of compression tests. They reported an exponential association between porosity of octahedral porous structures in Ti-6Al-4V and experimental fracture load [53]. Sercombe et al. [54] coupled in situ compression testing X-ray micro tomography (XMT) with finite element analysis (FEA) to investigate the failure mechanisms of high strength and stiffness to weight Ti-6Al-4V scaffolds produced using SLM. It is

Additive manufacturing of cp-Ti, Ti-6Al-4V and Ti2448

313

Top view

Max. principal stress (MPa) 10

(A)

(B) 0 −4.4

Side view 2mm

(C)

(D)

Fig. 3.6.5 XMT and corresponding FEA of the central layer of unit cells for the 10% solid fraction scaffold. This scaffold has been loaded well beyond the peak load and exhibits multiple fractures of the horizontal arms (arrowed). The top view is shown in (A) and (B) and the side view in (C) and (D). It is apparent that the failure sites are strongly correlated with the high stress regions in the FEA, which are also indicated with arrows [54].

apparent that the failure of these structures first occurs in the struts that carry the tensile load. Further, failure occurred at the sites that FEA predicted to have high localized stress due to poor build quality (Fig. 3.6.5 [54]). Liu et al. [46] have investigated the effect of the processing parameters using five different laser scan speeds from 500 to 1500 mm/s on the quality and mechanical properties of a biomedical Ti2448 alloy scaffold fabricated by SLM. Optimal manufacturing parameters were then determined through analyzing the pore distribution, geometrical accuracy, and mechanical properties of the produced components (Fig. 3.6.6 [46]). Scaffolds parts with near full density of solid strut (>99%) were obtained at a laser power of 175 W and a scan speed of between 750 and 1000 mm/s. Using these optimal processing conditions, the strength of the scaffold reached 51 MPa at a scaffold density of 99%. Analysis of the fracture surface revealed that the main reason for strut failure was the weakness of struts caused by the variable thickness of struts, pores, and unmelted powders inside solid strut parts. The failure appears to occur in the horizontal arms that carry the tensile load (Fig. 3.6.6C and D [46]).

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500 mm/s 750 mm/s 1000 mm/s 1250 mm/s 1500 mm/s

1200 180 150 120 90 60 30 0

Count

900 600 300

40 50 60 70 80 90

Compressive stress (Mpa)

1500 50 40 30 500 mm/s 750 mm/s 1000 mm/s 1250 mm/s 1500 mm/s

20 10

0 0

20

(A)

40

60 80 100 120 140 160 Pore size (mm)

0

0.02

(B)

0.06

0.10

0.14

0.18

Strain (mm/mm)

Y

X

(C)

3mm

(D)

Build direction (Z axis)

Fig. 3.6.6 Structural feature and compressive performance of SLM-processed Ti-24Nb-4Zr-8Sn scaffolds: (A) the distribution of pore sizes and amounts in solid strut of scaffolds with different laser scan speeds, (B) typical compressive stress-strain curves with the arrows indicating the location of the first strut failure, and visualization of the deformed scaffolds: (C) the middle plane of cells showing the location of the failure (black circle) and (D) a higher magnification view of the crack [46].

3.6.3.2.2 Electron beam melting of porous titanium structures Studies on EBM of porous titanium structures have largely focused on Ti-6Al-4V, although work on Ti2448 has also been reported. Li et al. [55] compared the compression fatigue behavior of EBM-produced Ti-6Al-4V mesh arrays with high levels of porosity ( 60%–85%). The fatigue lives of the EBM-produced Ti-6Al-4V mesh arrays were mainly determined by uniform deformation through the entire specimens while their failures were characterized by rapid strain accumulation and a severe crush band at an angle of 45 degrees to the cyclic loading direction (Fig. 3.6.7 [55]). The underlying mechanism of the fatigue failure appeared to be an interaction between

Additive manufacturing of cp-Ti, Ti-6Al-4V and Ti2448

315

45°

1

2

3

3

4

0.05

Accumulated strain

2 0.00 1

−0.05

3

−0.10

4

−0.15 −0.20 100

101

102

103

104

105

Cycles Fig. 3.6.7 Typical variation of the accumulated strain with cycle number of the EBM-produced Ti-6Al-4V mesh array. Macro observations are shown for (1)–(4) [55].

cyclic ratcheting and fatigue crack initiation and propagation, with the former playing a dominant role in fatigue life. Herna´ndez-Nava et al. [56] investigated the effect of defects on the mechanical response of EBM-produced Ti-6Al-4V cubic lattice structures. Internal pores for struts aligned with the build direction were found around the edges of the solid structures in regions that seem to be associated with the electron beam scan pattern. Although struts normal to the build direction showed more significant defects, their redundancy meant that they did not compromise the compressive performance in the build direction.

3.6.3.2.3 EBM versus SLM porous structures Liu et al. [57] compared the difference in the microstructure, defects, and mechanical behavior of porous structures from a beta-type Ti2448 alloy manufactured by EBM and SLM. As seen in Fig. 3.6.8 [57], the relative density calculated based on micro-CT 3D data was 99.94% and 99.63% for the EBM- and SLM-produced specimens, respectively. The SLM-produced samples also exhibited a smoother strut

Titanium in Medical and Dental Applications

Building direction

316

Count

100

(e)

EBM SLM

10

1

20

40

60 80 100 120 140 Defects size, µm

Fig. 3.6.8 The micro-CT reconstructed images showing the strut outside surface of the (A) EBM- and (B) SLM-produced Ti2448 samples, the defects inside the solid struts of the (C) EBM- and (D) SLM-produced Ti2448 samples, and (E) the size and count distribution of the defects inside the samples as a function of equivalent diameter [57].

surface than the EBM-fabricated counterparts. As can be seen from the location and distribution of the defects (marked in red) inside the solid struts in Fig. 3.6.8C and D [57], nearly all the defects in the EBM-produced samples were spherical in shape. In contrast, those produced via SLM had a more irregular shape, including some that were conical. It was also noted that the number of defects inside the EBM samples was less than that of SLM samples (Fig. 3.6.8E [57]); this was especially true for pores with an equivalent diameter (EqDiameter) of 30–90 μm. The EqDiameter indicates

Additive manufacturing of cp-Ti, Ti-6Al-4V and Ti2448

317

the equivalent diameter of the spheres that have the same volume as the measured defects. In the as-fabricated state, the EBM-produced porous Ti2448 samples had a higher Young’s modulus (1.34  0.04 GPa) than the SLM-produced ones (0.95  0.05 GPa) (Fig. 3.6.9A [57]). This was attributed to the presence of the higher modulus α phase in the microstructure of EBM-produced samples—a result of the part bed being kept at an elevated temperature during the build, causing both a slower cooling rate and effective age hardening [57]. After annealing, the Young’s modulus of the EBM samples had decreased to 1.04  0.04 GPa, which was very close to that of the SLM samples (1.09  0.03 GPa). In order to remove the effect of the presence of different phases, the fatigue performance of the scaffolds was measured on annealed samples, which contained only a single β phase (Fig. 3.6.9B [57]). The fatigue strength of the annealed EBM or SLM samples was clearly sensitive to the stress amplitudes. At the lower stress levels, the fatigue behavior of the meshes was mainly determined by the cyclic ratcheting and surface properties of the struts, resulting in similar properties for both manufacturing processes. However, at higher stress levels, the crack initiation and propagation from the pores tended to occur; therefore the SLM samples, which contain a higher number of defects, had a lower and more variable fatigue life. Liu et al. [58] have also studied the structural features (i.e., microstructure, strut surfaces, and defects) and the mechanical behavior of EBM-fabricated beta-type Ti2448 porous components with 70% porosity. Lower electron beam scan speeds led to more input energy, thereby producing stronger struts with fewer defects. This resulted in better mechanical properties. As seen in Table 3.6.4, EBM-processed Ti2448 porous components have at least twice the strength-to-modulus ratio of Ti-6Al-4V porous components produced using the same structure and at the same porosity. These excellent properties were attributed to the precipitation of the α phase at β grain boundaries due to high-temperature preheating in the EBM process [58]. Such nontoxic Ti2448 porous components with very high strength-to-modulus ratios are highly attractive for biomedical applications. 13 Compressive stress, MPa

Young's modulus, GPa

1.4 1.2 1.0 0.8 0.6 0.4 0.2 0.0

(A)

EBM

SLM

EBM SLM annealed annealed

12 11 10 9 8 7 6

Anealed EBM-produced sample Anealed SLM-produced sample

5

(B)

106 105 Number of cycles to failure

Fig. 3.6.9 (A) The Young’s modulus for EBM- or SLM-produced porous Ti2448 samples, and (B) the fatigue of samples annealed at 750°C for 1 h [57].

318

Table 3.6.4 Compressive mechanical properties, Vickers hardness (HV), and phase constituents of EBM- and SLM-manufactured titanium materials Method

Phase constituent

HV

E (GPa)

σ max (MPa)

σ max/E (×1023)

Reference

Ti2448 (solid) Ti2448 (solid) Ti2448 (70% porosity) Ti6Al4V (70% porosity) cp-Ti (37 porosity) Ti-8.35 vol% TiB (37 porosity)

EBM SLM EBM EBM SLM SLM

β β α+β α+β α + α0 α+β

250 228  6 280  5 331  14 261  13 [50] 402  7 [64]

… 53 1a 0.7 0.1 1 13 3 25 2

… 66518a 35 2b 20b 235 52 256 4

… 12.5 50 20 18.1 10.2

[59] [27] [58] [60] [48] [48]

Young’s modulus E, ultimate strength σ max, and strength-to-modulus ratio σ max/E [58]. a In tension. b First strut failure strength.

Titanium in Medical and Dental Applications

Material

Additive manufacturing of cp-Ti, Ti-6Al-4V and Ti2448

319

Stress/Yield stress

1

Ti2448 samples with 75.0% porosity

0.1

A(67.9%) B(72.5%) C(75.0%) D(77.4%) E(79.5%) F(91.2%) Ti-6AI-4V(75%) 1.0×104

1.0×105

Fig. 3.6.10 The normalized S-N curves of the EBM-produced porous Ti2448 and Ti-6Al-4V specimens with different porosity [61].

Ti-6AI-4V samples with 75% porosity 1.0×106

1.0×107

Number of cycles

In order to further understand the mechanical property-porosity relationships as well as the fatigue crack deflection behavior in Ti2448 porous samples, Liu et al. [61] investigated the influence of porosity variation on the mechanical properties of the β-type Ti2448 alloy porous samples, in terms of Young’s modulus, superelastic properties, strength, and fatigue properties. For Ti2448 samples, the fatigue life was dominated by the porosity and applied stress level; the fatigue life decreased with increasing porosity for both low and high stress levels, as reported in [62]. Unsurprisingly, the high-porosity samples (such as the group with 91.2% porosity) resulted in much lower fatigue strengths than the low-porosity groups. However, their normalized fatigue strength (Fig. 3.6.10 [61]) was almost independent of porosity level, and was also much higher than that measured for the Ti-6Al-4V samples with 75% porosity. This was thought to be a result of the superelastic property of this material and the larger plastic zone ahead of the fatigue crack tip. For the same fatigue strength, the Young’s modulus of Ti2448 porous samples is only half that of Ti-6Al-4V porous samples.

3.6.4

Biocompatibility of AM porous Ti

Good biocompatibility is a basic requirement for an implant material. Metallic implants remain in long-term contact with bodily fluids and tissues, which may lead to corrosion and the release of alloying elements into the body that may cause adverse effects [63–65]. As such, the biocompatibility of SLM- and EBM-produced titanium alloys must be investigated prior to any potential clinical application. For porous structures, an interconnected porous structure along with observed osteoblast cellular activity (cell proliferation, cytoplasmic extensions, synthesis of intracellular and extracellular proteins, differentiation, and mineralization) are beneficial. The interconnected porous structure also provides a pathway for adequate supply of nutrients

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and oxygen to cells and tissue. This prevents cell death associated with necrosis and/or hypoxia. Kumar et al. [66] tested cell-derived decellularized extracellular matrix (dECM) for porous Ti-6Al-4V scaffolds in vitro. The bioactive factors are found in the extracellular matrix, which may improve the cell functionality growth on Ti-6Al-4V scaffolds. Wang et al. [67] compared the biocompatibility, including cytocompatibility, haemocompatibility, skin irritation, and skin sensitivity of Ti-6Al-4V fabricated by EBM and SLM. Both the EBM- and SLM-produced Ti-6Al-4V parts exhibited good cytobiocompatibility and good hemocompatibility. The EBM- and SLM-produced Ti-6Al-4V samples showed no dermal irritation when exposed to rabbits. In a delayed hypersensitivity test, no allergic skin reaction from the EBM or the SLM Ti6Al4V was observed in guinea pigs. Based on these results it may be concluded that Ti6Al4V fabricated by EBM and SLM are cytobiocompatible, hemocompatible, nonirritant, and nonsensitizing materials. The in vitro biological activity of the EBM Ti2448 alloy mesh structures was also studied in terms of bioactivity, osteoblast cell attachment, proliferation, differentiation, and mineralization [68]. The results showed that the Ti2448 alloy mesh structures have high cell viability and good biocompatibility. The results also showed that 3D-printed Ti2448 alloy mesh structures could aid in the bone healing process by providing a favorable osteogenic microenvironment for tissue in-growth.

3.6.5

Summary

The combination of high strength, low weight, excellent corrosion resistance, and biocompatibility combined with a relatively low modulus make titanium and titanium alloys widely used in load-bearing biomedical applications. However, despite the lower modulus, conventional titanium alloys have a modulus at least five times that of bone, which can lead to stress shielding. As a result, there has been a significant drive to develop new beta-titanium alloys that have a much lower modulus. One such alloy, Ti2448, is proving particularly promising due to its combination of very low modulus (450 MPa, a yield strength of >300 MPa, and an elastic modulus of
110 GPa with an average hardness of >150 HV [10]. With the addition of aluminum and/or vanadium in CpTi, the phases present can vary. Vanadium is a β-phase stabilizer while aluminum is a α-phase stabilizer; some other elements have no effect on phase stabilization and are considered neutral. Graphical representation of how α or β-phase stabilizers or neutral alloying elements influence the phase diagram is shown in Fig. 3.7.1 at varying compositions and temperatures [11].

Titanium in Medical and Dental Applications. https://doi.org/10.1016/B978-0-12-812456-7.00015-9 © 2018 Elsevier Inc. All rights reserved.

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α

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Fig. 3.7.1 Phase diagrams as a function of alpha, beta or neutral stabilizing constituent presence [11]. Reprinted with permission from Elsevier License #4183260336823. Original article may be viewed at: https://doi.org/10.1016/j.msec.2016.10.025.

The most predominately used titanium alloy is Ti-6Al-4V (Ti64), having an average ultimate tensile strength of >1000 MPa, an average yield strength of >900 MPa, and an elastic modulus of
110 GPa with an average hardness of >350 HV [10]. Ti64 is bioinert and, when compared to pure titanium, exhibits higher fatigue resistance due to the (α + β) microstructure. One issue with Ti64 is of potential vanadium ion leaching, which can be toxic to the body. However, among millions of Ti64 implants currently in use, toxicity due to V ion release is rarely reported in the literature. Ti64 is also a soft metal and shows poor wear resistance. This is a concern for articulating surfaces of implants for applications such as balls or acetabular cups of hip implants. For this reason, in total joint replacement applications, either the articulating surface is coated with a ceramic, a wear-resistant medical-grade polyethylene, or an ultrahigh molecular weight polyethylene (UHMWPE) to reduce metal ion release in vivo.

3.7.1.2 Why additive manufacturing? Additive manufacturing (AM) or three-dimensional (3D) printing is a layer-by-layer fabrication process where the part of choice is built from a 3D computer-aided design (CAD) file. The file is first sliced along the z-axis in a virtual environment, and then for each slice a machine-specific tool path is generated. Several fabrication techniques exist for metals, but we will only focus on three of the most popular ones for biomedical applications: (1) electro-optical systems (EOS)-selective laser melting (SLM), (2) Arcam-electron beam melting (EBM), and (3) laser-engineered net shaping (LENS™). AM has become very popular due to its build-on-demand attributes and fixed cost toward low-volume production when compared to many traditional fabrication methods. Also, the ability to build geometrically intricate parts without any significant additional costs or manufacturing of patient-specific devices makes AM the exciting new kid in town for orthopedic and dental implant manufacturing practices. For orthopedic implants, one of the major factors that determines the implants’ in vivo longevity is osseointegration, that is, how well the bone forming cells (osteoblasts) can adhere to the surface of the implant. Osseointegration depends on surface properties such as wettability, surface energy, chemistry, topography, surface charge,

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and porosity [12]. As the demand for joint replacements is projected to increase by 174% for hip and 673% for knee implants from 2009 to 2030 [13], the interest in innovative implants with improved osseointegration ability for longer in vivo life is growing. Among other options, AM techniques hold a significant promise to accomplish this goal in addition to the ease of manufacturing complex devices on demand. Traditionally, the metals of choice for dental implantation are cobalt-chromium and titanium alloys. These alloys are usually conventionally manufactured with investment casting, that is, the lost wax technique [14,15]. In the lost wax technique, wax is burned out by the molten metal, allowed to solidify and part finishing is performed to match design tolerances. Many issues have arisen when dealing with the lost wax technique, such as casting defects, wax distortion, and longer processing time. This makes it somewhat inconvenient for patients because of multiple visits to the dentist’s office for the same treatment [16]. Application of AM-based techniques is becoming popular due to minimizing the number of visits to dentist offices for on-demand, defect-matched, patient-specific implants.

3.7.2

Additive manufacturing processes

3.7.2.1 EOS-SLM EOS is a powder bed-based metallic AM system called SLM. The SLM process starts by selecting the base metal of choice. Typically called the “substrate” in the field, this base metal is where the structure of choice is “printed on,” and may or may not be required to be removed during post manufacturing. The metal powder of choice for part fabrication is spread over the substrate at a predetermined layer thickness that allows for maximum efficiency and density of the part during the layer-by-layer build process. Following that, a high-powered laser is focused onto the metal powder layer, creating a melt pool. The laser continues through its path and fuses the metal powder particles together to produce the cross-sectional design in the plane of the metal powder. A sequential layer is then spread over the existing pattern and the process is continued. The melt pool is protected from oxidation by blowing an inert gas over it, similar to TIG welding. An advantage to SLM is that the unmelted powder in the powder bed also serves as a support material. This, however, confines the part size to the size of the powder bed and not the structural dimensions possible through CAD software. EOS’s SLM technique is widely used in various applications including aerospace, biomedical devices, and the automotive and machine-tool industries, to name a few. In today’s metal-based AM environment, it is one of the most popular systems used worldwide.

3.7.2.1.1 EOS-SLM of beta-type Ti-15Mo-5Zr-3Al alloy Current research using SLM has produced components where texture control of the surface can be achieved. Such research has resulted in being able to reduce the Young’s modulus of biocompatible implant materials such as Ti-15Mo-5Zr-3Al to below 70 GPa from the reported >80 GPa [18], which is lower than the previously reported Ti64 value of >110 GPa. Implants with such mechanical properties have

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the potential to minimize the unintended phenomena of bone resorption or stress shielding of the bone surrounding the implant due to a mismatch in Young’s moduli between the cortical bone and the implant [19]. Human cortical bone has shown a typical modulus between 10 and 30 GPa [17]. Such proof-of-concept work was carried out to utilize the low modulus of β-Ti. Ti-15Mo-5Zr-3Al has only the β-Ti phase and has been allowed by the International Organization for Standardization (ISO) for use as a biomaterial (ISO 5832-14) since 2007. The objective of the work was to decrease the modulus by using single crystal orientation. Past work has proven that the modulus depends strongly on crystal orientation in the Ti-15Mo-5Zr-3Al alloy [20]. Specifically, for this alloy, these values are
120 GPa in the h111i directions and
44 GPa in the h001i directions in its bcc crystal structure. Dense samples were produced with preferred crystal orientation by means of tuning the SLM parameters. Two scanning strategies used were referred to as “Scan Strategy X” and “Scan Strategy XY,” where the first was a bidirectional (zigzag) path in the x-axis and the second was bidirectional with 90-degree rotations between layers. Such results are similar to what has been previously reported in materials with bcc and fcc structures [21–23], showing a strong h001i and h011i orientation in the x and z directions with Scan Strategy X and a strong h001i orientation along the x, y, and z directions with Scan Strategy XY. These results are also shown in Fig. 3.7.2.

3.7.2.1.2 SLM of porous implants with immobilized silver particles SLM-produced implants can be processed further not only to increase their biocompatibility, but also to incorporate antibacterial properties through silver (Ag) introduction. Silver nanoparticle inclusion into implants has been shown to behave as a powerful ScanStrategy X

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Fig. 3.7.2 Inverse pole figure (IPF) images taken in the three orthogonal planes. (A), (C) Inverse pole figure (IPF) images taken in the three orthogonal planes. (B), (D) {001} and {011} pole figures of the products measured in the y-z plane [17]. Reprinted with free permission. Original article may be viewed at: https://doi.org/10.1016/j. scriptamat.2016.12.038, https://creativecommons.org/licenses/by/4.0/.

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antimicrobial agent against many bacteria, including methicillin-resistant Staphylococcus aureus (MRSA), which is a bacterium that is known to be resistant to many antibiotics [25–30]. One such mechanism of silver’s ability to produce these effects is its ability to damage the membrane of the bacterial cells and produce reactive oxygen species [31]. Being able to produce a multifunctional orthopedic implant would most benefit patients with compromised bone metabolism and immune systems. An example would be patients with malignant bone tumors, in which case surgery would salvage large limbs [32–35]. Further driving this work is the fact that by tailoring geometrical parameters in porous implants, such as size and shape of the pores, tissue regeneration can be improved toward enhanced biological fixation of the implant [36–39]. The objective of this work was to produce porous and rationally designed metallic implants with antimicrobial properties to prevent implant-associated infections. Porous implants were fabricated using Ti64 and further processed to grow an oxide layer on the surface through plasma electrolytic oxidation (PEO) in a Ca/P-based electrolyte followed by embedding the coating with silver nanoparticles. This approach resulted in a bioactive surface with interconnected pores for bone tissue integration. Pore interconnectivity was shown to exceptionally improve implant osseointegration [40–43]. The SLM PEO + Ag implant was compared to a PEO + Ag implant with the only varying factor being SLM production and non-SLM production. Upon performing silver release studies up to 28 days, it was found that the SLM-produced implant showed a higher release (4.35 times) than the non-SLM produced implant. This is due to the inherent increased surface area due to the layer-by-layer deposition in SLM. It is reported that a minimum of 50 ppb of silver released continuously confined in nanoscale is enough to achieve antimicrobial characteristics in vivo [44]. The PEO + Ag implant exhibited roughly 100 ppb while the SLM PEO + Ag implant was significantly higher, as shown in Fig. 3.7.3.

Fig. 3.7.3 Cumulative silver ion release profiles (n ¼ 3) of implants measured by ICP-OES [24]. Reprinted with permission from Elsevier License #4183261002674. Original article may be viewed at: https://doi.org/10.1016/j.biomaterials. 2017.02.030.

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3.7.2.2 Arcam-EBM EBM is another powder bed metal AM technique that can produce complex 3D parts in which a high intensity electron beam is used as the heating source to melt metal powders. The Arcam instrument uses CAD files and allows for the fabrication of multiple parts in the same build. CAD files are uploaded into the instrument. Then, software slices the part along the z-direction into cross sections to be built in a layer-by-layer process from metal powders. Powder is poured into the Arcam chamber, then funneled and smoothed over a recessed stage by raking, which continues to drop in the z-direction depending on the desired slice thickness for the cross sections. The powder is then partially heated by the stage before it is fused by the focused electron beam. After the fusion of the layer is complete, the stage drops once again and is raked, reheated, and fused, creating the layer-by-layer process. The entire process is performed under high levels of vacuum in the chamber; this is done to prevent the molten material from reacting with oxygen. One of the advantages of EBM over SLM is that, due to the heating of the powder bed, residual stresses are minimized from the top layer to the bottom layer. Also, the electron beam is controlled by an electromagnetic lens that allows for faster control of beam size and translation when compared to the mechanical mirror system associated with SLM; this feature can be translated to faster builds using EBM. Arcam’s EBM has seen applications in medical implants, aerospace, and defense as well as various other manufacturing and product development industries.

3.7.2.2.1 Arcam-EBM of customized Ti64 dental implants Previous work has been performed on producing EBM dental implants of Ti64 that have shown improved mechanical properties compared to implants produced via metal casting [46]. Such results are utilized toward customized roots of dental implants manufactured via EBM technology [45]. The objective of this research was to evaluate important parameters such as surface microstructure, topography, chemical composition, and wettability toward customized roots of dental implants via EBM. Such parameters are important for osseointegration. Fig. 3.7.4 shows an EBM manufactured sample, where Fig. 3.7.4A is of the.stl design file and Fig. 3.7.4B is a μCT scan of the fabricated implant revealing the core region plane and showing a dense part with an isolated pore. By examining the surface topography, it was reported that better sintering of the metal powder was observed, shown in Fig. 3.7.4C, along the build plane. However, the lateral surface was rougher and displayed incompletely sintered metal powder particles, shown in Figs. 3.7.4D–F. Surface roughness was calculated to be 0.682 and 3.398 μm for build and lateral surfaces, respectively, and was attained after applying a Gaussian filter. Upon characterizing the implant’s surface chemistry, it was reported that the elements and their respective weight percent were as follows: Ti
88.5%, Al
6.95%, V
3.38%, C
0.74%, O
0.014%, H
0.013%, Fe
0.25%, and N
0.035. When determining the contact angle for surface wettability, the angle was reported to be zero due to inherent surface roughness. It has been conveyed that increasing surface wettability, that is, a

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Fig. 3.7.4 (A) Designed. stl picture (B) Micro CT image of the root form implant taken at the mid-section longitudinal slice, showing dense internal structure and isolated internal porosity. SEM images (C) build surface showing the uniformly melted alloy particles and edge of lateral surface having partially melted alloy particles (D), (E) and (F) lateral surface showing partially melted alloy particles at different magnifications [45]. Reprinted with free permission. Original article may be viewed at: https://doi.org/10.1016/j.sjbs.2016.05.001, https://creativecommons.org/licenses/ by-nc-nd/4.0/.

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highly hydrophilic surface, increases the early-stage bone tissue integration or osseointegration [47]. Also, implants for biomedical applications exhibiting a higher degree of wettability can help in the adhesion of proteins, interstitial fluid, and blood macromolecules, which can further improve biological integration [48,49] and result in faster healing for patients. Moreover, commercially available implants are generic in geometry and may not offer the best implant-patient match [45]. Using AM-based technology, the defect site can be scanned and alterations can be performed to improve fit and match for a specific patient as desired in less time [50].

3.7.2.2.2 Arcam-EBM for trabecular titanium structures in orthopedic implants Oftentimes for articulating joint implants, the need for increased osseoingegration will require the need for improved surface topography that is most ideal for bone integration. Thus, this requires a more systematic build to produce a geometrically symmetric implant surface topography structure [51]. For example, hip implants would require increased integration of hard tissue into the implant. This is to withstand the amount of fatigue and load the implant will experience over its lifespan, primarily meaning the exterior surface of the acetabular cup. As mentioned earlier, porous implant surfaces improve bone integration, thus this research has brought the current design of incorporating a highly porous trabecular titanium (TT) onto the surface of the implant into the market of modifying existing implant surfaces. Also, the effort for producing such implant surface structure is fueled by the reported value of 600 μm diameter pores allowing for the fastest osteoblast growth [52]. Moreover, it has been reported that microtexture can improve bone tissue in-growth [53,54]. The word trabecula is Latin, meaning “small beam.” The honeycomb structure comprised of small beams can be seen in Fig. 3.7.5A, representing the 3D CAD image. A balance of pore size and shape and mechanical properties without overlooking biocompatibility and corrosion resistance is taken into consideration when designing and producing such structures.

Fig. 3.7.5 (A) cubic CAD 3D model made of cellular solid structures. (B) SEM image of the cellular solid structure as seen from the side with porosities in evidence [51]. Reprinted with permission from Elsevier License #4183270083563. Original article may be view at: https://doi.org/10.1016/j.jmbbm.2010.02.001.

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For characterization purposes, it was reported that two cylindrical scaffolds (10.5 mm diameter  12 mm height) with differing design pore diameter (650 and 1400 μm) were produced for compression testing; dog-bone flat samples were produced for tension testing. It is known that a minimum beam spot diameter threshold exists when producing and controlling a melt pool. This limitation can be seen by the arms and node visible in Fig. 3.7.5B. As seen, the pores do not appear as hexagonal and edges are rounded off. The mean pore area was reported to be 0.33  0.11 and 1.61  0.15 mm2, the mean equivalent pore diameter to be 0.64  0.11 and 1.43  0.07 mm, a porosity of 63% and 72%, and a hardness of 381.6  17.4 and 385  8.8 HV0.1 for the designed 650 and 1400 μm structure, respectively. It should be noted that the designed 650 μm structure is most appropriate for osteoblast cell growth, considering mean pore diameter. Tension tests were conveyed under three angle orientations with respect to the TT layered structure and the values were 4312  12.16, 4368.56  14.24, and 4754.36  17.68 MPa for orientation angles 90 degrees, 45 degrees, and 0 degrees, respectively. These elastic modulus values were far inferior to that of cortical bone (no stress shielding should be expected) and more comparable to that of cancellous bone. It can be seen (Fig. 3.7.6A) that the TT structure was implemented onto the acetabular cup and is observed in vivo (Fig. 3.7.6B) and should be expected to improve the early stage osseointegration of the implant. The reported work displayed that, via EBM technology, comparable porosity to that of spongy (cancellous
50%–80%) bone can be achieved. The pore size most suitable for improved osseointegration as found in the literature was comparable and had been achieved for already biocompatible Ti64 biomedical implants. Tension testing showed that detachment of the TT structure and bulk material was not observed.

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Fig. 3.7.6 (A) Acetabular cup in trabecular titanium (Delta TT, Lima-Lto, Villanova San Daniele del Friuli, Italy) for use in primary acetabular surgery. Note that the external microstructure surface in TT that imitates the morphology of the trabecular bone. (B) Postoperative radiograph showing the use of the Delta TT acetabular component in situ for 6 months. The bone-implant interface appears stable, with a slight reactive line in the superior area [51]. Reprinted with permission from Elsevier License #4183270083563. Original article may be viewed at: https://doi.org/10.1016/j.jmbbm.2010.02.001.

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3.7.2.3 Laser-engineered net shaping LENS™ is a laser-based AM technology that does not use a powder bed. In this technique, a chamber houses the build stage with its own x-y axis of translation while the laser head moves along the z-direction. Powder is delivered at the focal point of the laser into the controlled atmosphere chamber by an inert gas. The powder and laser create a melt pool on the surface of the substrate metal, which is then used for printing. Because LENS™ is not a powder-bed technique, parts fabricated from this technique are not restricted to a powder-bed size and can be up to several feet in length. Unlike the previously mentioned techniques, LENS™ provides the ability to produce parts with multiple materials such as compositionally graded dense or porous structures; such variations can be achieved while the part is being printed. Initial research toward the use of LENS™ technology for biomedical implants started with low-stiffness porous implants [55] with optimized pore geometry to enhance in vitro bone cell in-growth [56]. Further research focused on producing graded structures for load-bearing implant applications [57] and manufacturing hard coatings for articulating surfaces using zirconium [58] or TiO2 [59] on titanium. Also, other biocompatible materials such as CoCrMo and NiTi have been used to modify their effective modulus by inducing porosity to reduce stress shielding [60].

3.7.2.3.1 LENS™—Influence of porosity on Ti6Al4V’s mechanical properties and in vivo response By tailoring the porosity of LENS™-fabricated implants, previous work has shown that the elastic modulus of Ti64 implants can be reduced to 7 and 60GPa [61]. This is significantly lower than the >110 GPa seen in dense Ti64 implants and allows for better elastic modulus matching with that of the cortical bone, which can range from 10 to 30 GPa. The porosity of orthopedic implants improves tissue adhesion, vascularization, and growth. The objective of the work was to produce porous Ti64 implants in an attempt to enhance biological fixation to increase tissue in-growth and reduce the modulus mismatch. Seven and 12 mm cylindrical samples of different porosities, as shown in Fig. 3.7.7A, were produced for mechanical testing by varying the hatch distance between 0.762 and 1.27 mm to control the level of porosity. Bulk density was measured by geometric means and mass while apparent density was determined by the Archimedes’ principle to determine the fraction of open pores and closed pores. Samples for in vivo studies had densities of 75%, 89.3%, and 97.2% and were implanted bilaterally (distal femur) in male Sprague-Dawley rats up to 16 weeks. Samples produced with a hatch spacing of 0.762 mm displayed an average pore diameter of 60–700  20 μm, as shown in Fig. 3.7.7B, while a hatch spacing of 1.27 mm showed an average pore diameter of 300–1500 90 μm (Fig. 3.7.7C). Mechanical testing resulted in elastic moduli of 7–60 GPa. Samples having porosity of 23–32 vol% exhibited the closest moduli to that of cortical bone. In vivo samples for porosities of 25%, 10.7%, and 2.8% were harvested and tissue adhesion was determined. The tissue in-growth is apparent in the 75% dense Ti64 implant (Fig. 3.7.8) when compared to the lower porosity implants.

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CM

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Fig. 3.7.7 (A) Ti6Al4V samples fabricated by LENS™. Micrograph showing surface porosity and the following pore interconnectivity for, (B) hatch spacing of 0.762 mm; (C) hatch spacing 1.27 mm [61]. Reprinted with permission from Elsevier License #4183270526983. Original article may be viewed at: https://doi.org/10.1016/j.actbio.2009.11.011. 97.2 ± 0.6% (Control)

89.3 ± 0.5%

75.0 ± 0.8% Biological tissue

Fig. 3.7.8 LENS™ processed porous Ti64 implants after 16 weeks implantation in rat intramedullary defects [61]. Reprinted with permission from Elsevier License #4183270526983. Original article may be viewed at: https://doi.org/10.1016/j.actbio.2009.11.011.

Through varying processing parameters during the LENS™ technique, visible signs of increased host-tissue integration into the implant were observed.

3.7.2.3.2 LENS™—In vivo response of laser-processed porous titanium implants for load-bearing applications In the following work, porous titanium implants were produced with approximately 25 vol% porosity by LENS™ [62]. The objective was to further understand the role of porosity on titanium implants with and without nanotexturing by TiO2 nanotube formation on the surface via electrochemical anodization. It was hypothesized that pore interconnectivity produced via the LENS™ technique would increase hard-tissue integration when compared to dense implants. Cylindrical rods of 3.0 mm diameter

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Fig. 3.7.9 (A) LENS™ processed porous Ti samples with 25% porosity. (B) SEM image of the porous surface nature of LENS™-processed porous sample. (C) Low magnification SEM image of porous Ti implant with fabrication of nanotubes (D) with diameter 105  30 nm and length 375  35 nm using anodization method [62]. Reprinted with permission from Springer License #4183290317785. Original article may be viewed at: https://doi.org/10.1007/s10439-016-1673-8.

were produced (Fig. 3.7.9A) and surface porosity was achieved by partially melting the metal powder with low laser power and parameter optimizations such as hatch spacing, laser speed, and layer thickness. Samples were ground to produce a slightly smoother surface (Fig. 3.7.9B) to prevent breaking of the distal femur of the Sprague-Dawley male rats when implanting. Samples needing TiO2 surface modification were carried out by an electrochemical anodization method in 1% hydrofluoric acid. Bilateral distal femur implantation was performed on rats for 4-week and 10-week intervals. Final implants for in vivo use can be seen in Fig. 3.7.9A. Push-out testing was performed to measure the interfacial shear modulus of the implant and host tissue. CT scanning, histology, and SEM characterization were also conducted. By SEM characterization, it was reported that volume porosity was 25% and pore size was in the range of 200–300 μm. Anodization produced nanotubes with lengths of 375  35 nm and diameter of 105  30 nm (Fig. 3.7.9D). CT images showed that no defects or gaps were present and good bonding was observed, even better for the porous samples where exceptional osseointegration was seen by high resolution CT scans, as shown in

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Fig. 3.7.10. Push-out data revealed that the shear modulus for the dense Ti, LENS™porous Ti, and LENS™-porous Ti-NT were 14.87  2.63, 25.82  1.94, and 29.38  2.52 MPa, respectively, after 4 weeks. Histological characterization revealed that as early as 4 weeks, osteoid-like new bone formation had occurred for porous implants. These are the orange-red regions in Fig. 3.7.11 while the greenish region represents mineralized bone; the nuclei are represented by the bluish-black regions. Finally, SEM micrographs of the histology samples showed a decrease in the gap at the host tissue implant location by porosity alone and a negligible gap with the porous-nanotube implant (Fig. 3.7.12). In summary, it was found that the LENS™ technique can mimic bone properties and increase early stage host tissue and implant integration. Cell in-growth into pores was achieved and interfacial bonding strength was increased, reported as early as 4 weeks. It was concluded that LENS™-produced implants, in the presence of or without surface modification, increase early-stage osseointegration. Also, defect healing occurs with TiO2 present on the surface of the implant, which has been shown to increase the interfacial bond between the host tissue and the implant [62].

Fig. 3.7.10 High-resolution micro CT images showing good interfacial bonding between the porous implants with the tissue along with the bone in-growth between the pores [62]. Reprinted with permission from Springer License #4183290317785. Original article may be viewed at: https://doi.org/10.1007/s10439-016-1673-8.

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Fig. 3.7.11 Photomicrograph showing the histology images after 4 weeks (A, B, C) and 10 weeks (D, E, F) where signs of osteoid-like new bone formation could be seen in orange/red color. Modified Masson Goldner’s trichrome staining method was used [62]. Reprinted with permission from Springer License #4183290317785. Original article may be viewed at: DOI: https://doi.org/10.1007/s10439-016-1673-8.

Implant

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Fig. 3.7.12 SEM images of stained 4 weeks dense Ti, porous Ti, and porous Ti-NT samples from left to right, respectively [62]. Reprinted with permission from Springer License #4183290317785. Original article may be viewed at: https://doi.org/10.1007/s10439-016-1673-8.

3.7.3

Challenges and future trends

One common challenge among the three fabrication techniques is the beam spot diameter. A minimum threshold is apparent when attempting micron or near-micron scale features in the build. The melt pool becomes hard to control and producing a precise “bead” that can replicate the desired architecture at that scale becomes increasingly hard.

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A patient’s habits and health condition can also contribute to some challenges in having biomedical implants not reach their intended life span. For example, patients with diabetes, smoking habits, and metabolic disease increase the possibility of implant failure when compared to healthy individuals. Infections are always of big concern. Moreover, improper bone bonding may also be observed due to overloading, such as an athlete resuming his/her natural activities. Ti64 is suitably biocompatible but issues arise in high load-bearing articulating surfaces. This is due to the possibility of having the cytotoxic elements aluminum and vanadium leaching into the body and causing severe issues [63]. Attempts at addressing these possible leaching issues have included producing new titanium alloys that consist of noncytotoxic elements. However, most are still in the clinicopathological stage and recent studies have shown that the elastic modulus of such titanium alloys has increased [64]. As known, this can cause the host tissue to become idle due to implant stress shielding and alternatively causing implant failure. Geometrical issues arise when balancing porosity and roughness to increase osseointegration with mechanical properties [65]. When designing a geometrically porous structure, even small variations in structural geometry can cause structural integrity failure by means of producing a stress concentration point. This must be addressed and balanced with an optimized geometry that will be free of stress concentrations while optimizing osseointegration. Finally, increasing surface implant complexity can have its drawbacks. Potential failures can arise from porous layer detachment, in addition to corrosion and implant loosening [65]. Porous layer detachments have been seen in hydroxyapatite (HA)coated implants. Cracking, scratches, or flaking may occur after surgery while under in vivo loading. Such failures could also accelerate the implant failure by allowing for exposed metal to corrode as well as allowing the release of metal ions [66]. The future for additive manufacturing of biomedical metal implants is not too far away. With increasing public interest and acceptance of modern-day implant technology, AM of metallic implants will only see an increase in magnitude. The demand for being able to efficiently produce unique and customized implants—and to do so cost effectively—will increase. In recent years, consumer interest in downloadable software for 3D printing designs increased due to decreased costs and expanding uses [67,68]. The potential for having a patient-specific implant ready just hours after an MRI or CT scan is also a possibility. Typically, what is seen in the orthopedic field is that standard implants may not fit well for some patients due to anatomical complexity or allergies. In those cases, surgeons need to improvise to allow the proper fit of the implant [69]. It is anticipated that in a decade, AM or 3D printing of Ti-based implants will be normal practice both in dentistry and orthopedics.

Acknowledgments Authors would like to acknowledge financial support from the National Institute of Arthritis and Musculoskeletal and Skin Diseases (NIAMS) of the National Institutes of Health under award number R01 AR067306-01A1. The content is solely the responsibility of the authors and does not necessarily represent the official views of the National Institutes of Health.

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Titanium spinal-fixation implants

4.1

M. Niinomi Tohoku University, Sendai, Japan, Osaka University, Osaka, Japan, Nagoya University, Nagoya, Japan, Meijyo University, Nagoya, Japan

4.1.1

Introduction

Owing to their superior strength and toughness, metallic materials for biomedical applications (referred to as biometals), including Ti (titanium)-based alloys, Co (cobalt)-based alloys, and stainless steel (mostly SUS 316L) are still very important, especially with regard to implants used under cyclic loading or high loading conditions. Among such biometals, Ti alloys receive much attention because their biocompatibility is highly superior to that of SUS 316L and Co-based alloys. CP-Ti (commercially pure) and Ti-6Al-4V ELI (extra interstitials) (T-64) are representative Ti-based biometals. However, these biometals have problems concerning their applications for implants. The strength of CP-Ti is not enough for load-bearing implants compared with that of other biometals such as Co-based alloys or Ti-based alloys [1]. However, Ti-64 shows a good balance of mechanical properties but contains V and Al, which have been designated as harmful elements to living tissue [2–4]. There is a significant difference between the rigidity, namely the Young’s modulus of the implant made of the biometal, and that of the cortical human bone; this can cause problems when using biometals for hard-tissue replacement [5]. Such a difference in the rigidity has a high possibility for causing the stress-shielding phenomena; this can reduce loads that are essential for the bone-tissue health, then cause a resorption of the bone. These also lower the quality of the bone surrounding the implant [6,7]. With regard to achieving low rigidity, β-type Ti alloys with a single β phase having a body-centered cubic (bcc) structure are advantageous as compared with α-type Ti alloys with a single α phase having a hexagonal close-packed (hcp) structure, and (α + β)-type Ti alloys having α and β phases because the atomic density of the β phase is less compared with that of the α phase. Therefore, many β-type Ti alloys with low rigidity and nontoxic elements have been developed as biometals [8]. Such Ti alloys with low rigidity have the potential to be used for various types of biometals for hard-tissue (bone) substitution implants. The rods used in spinal-fixation implants, with low rigidity, should be sufficiently flexible to prevent adjacent segmental diseases. Therefore, Ti alloys with low rigidity are considered suitable for use in the rods of spinal-fixation implants. Recently, a further concept regarding the spinal-fixation rods has been proposed. This concept develops β-type Ti alloys having changeable rigidity. These alloys could Titanium in Medical and Dental Applications. https://doi.org/10.1016/B978-0-12-812456-7.00016-0 © 2018 Elsevier Inc. All rights reserved.

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satisfy the demands of both surgeons and patients, and be used for spinal-fixation implants [9]. Ti and its alloys are also used in cages and wires for spinal-fixation implants. Such Ti alloys, which are targeted for use in the rod, cage, and wire of spinal-fixation implants, will be described in the next chapter.

4.1.2

Requirements for spinal-fixation rods

Biometals should consist of nontoxic and nonallergic elements. They should also possess high mechanical properties and resistance to corrosion to inhibit the dissolution of metallic elements, and a low rigidity close to that of bone. The demands mentioned above also apply to spinal-fixation rods. With regard to the high mechanical properties of spinal-fixation rods, the fatigue strength is a very important factor because tight bone fusion is generally targeted. This is because the spinal-fixation rod (implant) must maintain a strength above that of bone until complete fusion is achieved, as schematically shown in Fig. 4.1.1 [10]. Furthermore, spinal-fixation devices are generally composed of rods, plugs, and screws as schematically shown in Fig. 4.1.2 [11]. In particular, the rods are bent when they are manually treated by surgeons within the tiny space inside a patients’ body to achieve the in situ spine profile [9]. The contoured shape of the spinal-fixation rod should be maintained. Therefore, reverse bending, known as springback, of the bent rod should be prevented. With regard to implant rods, the degree of springback should be low to facilitate their handling during operations. Both the strength and rigidity of the spinal-fixation rod is thought to be related to the degree of springback. If two spinal-fixation rods with the same strength but differing rigidity are utilized, the spinal-fixation rod with the lower rigidity will exhibit larger springback, as schematically shown in Fig. 4.1.3 [11]. Therefore, with regard to the patients’ needs, low rigidity is required to inhibit stress shielding whereas to facilitate surgery, a high Young’s modulus is required to prevent springback. To satisfy these conflicting demands simultaneously, increasing the rigidity of the bent parts of the spinal-fixation rod via deformation at room temperature should be possible while the rigidity of

Strength

Implant

Complete bone fixation Insufficient bne fixation

Grafted bone

Failure of implant Time

Spinal fusion

Fig. 4.1.1 Roll of implants ¼ strong fixation up to complete bone fusion.

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349

Plug

Rod

Screw

Fig. 4.1.2 Schematic drawing of a spinal-fixation system consisting of rods, screws, and plugs.

Stress

High

Fig. 4.1.3 Relationship between Young’s modulus and springback.

Low

Young’s modulus Springback

Strain

Small Large Elastic recovered strain

the remainder of the spinal-fixation rod is allowed to remain unchanged at a low value [9]. Therefore, specifically, surgeons need materials with a high rigidity to inhibit springback during operations because of the limited operating space available within a patient’s body. However, with regard to the health of patients, materials with a low rigidity are required to inhibit the stress-shielding effect [6,7]. The spinal-fixation rods especially need to have low rigidity, excellent biocompatibility, and low springback. Therefore, the development of novel Ti alloys that offer excellent biocompatibility and a changeable rigidity is required. To accomplish this, increasing local rigidity to a high value at a certain part of the spinal-fixation rod through deformation at room temperature should be possible while the rigidity of the remainder of the spinal-fixation rod is allowed to remain unchanged at a lower value [9]. Moreover, three techniques are used for spinal fixation: spinal fusion, spinal stabilization, and conservative treatments as schematically shown in Fig. 4.1.4 [10]. When rods for spinal-fixation implants are fabricated using Ti alloys with low rigidity, they can be applied using spinal fusion or stabilization techniques.

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Risk for secondary fracture Low High

Operation treatment of spinal disease Spinal fusion

Target of low Young’s modulus titanium alloy Spinal stabilization Conservative treatment

Short

Long Period for complete bone fusion

Fig. 4.1.4 Schematic drawing of the risk for secondary fracture and the period for complete bone fusion in the operational treatment of spinal disease.

4.1.3

Advantages of Ti alloys with low rigidity for spinal-fixation implants

Animal tests have previously been carried out to investigate the effects of the rigidity of biometals on stress shielding and bone remodeling. For example, investigations on Japanese white rabbits were carried out using intramedullary-rod [12] and bone-plates implants [6]. These implants were manufactured using a Ti-29Nb-13Ta-4.6Zr (referred to as TNTZ) with a Young’s modulus of around 58 GPa. This is a β-type Ti alloy with a low rigidity intended for biomedical applications. In addition, SUS 316L with a Young’s modulus of around 161 GPa and a (α + β)-type Ti-64 with a Young’s modulus of around 108 GPa, which is the most widely used Ti alloy for biomedical applications, were also used to manufacture implants. The Young’s moduli of the implants were evaluated using three-point bending tests [12]. With regard to implanting the intramedullary rods and bone plates, both the lowest bone resorption and most excellent bone remodeling have been reported when the TNTZ parts were used. During the investigation on bone remodeling, bone plates manufactured from TNTZ, Ti-64, and SUS 316 were inserted in fracture models made in rabbit tibiae. Finally, only for the bone plate manufactured from the TNTZ, an increase in the diameter of the tibia and double-wall structure of the intramedullary bone tissue has been observed, as shown in Fig. 4.1.5 [6]. In this figure, the original bone, that is, the remaining old bone, coincides with the inner-wall structure of the bone whereas the newly formed part coincides with the outer wall of the bone structure. This bone remodeling is associated with the direct result of the use of a bone plate with low rigidity. Therefore, the low rigidity of the material is effective with regard to preventing bone resorption leading to excellent bone remodeling. A spinal-fixation implant with screws made of Ti-64 and rods made of TNTZ conducted with a solution treatment (referred to as TNTZ-ST) was implanted in sheep. Radiographs of the spines of the sheep with the implanted spinal-fixation implants were captured from the sagittal and coronal planes immediately following the surgery as well as 5 months later to evaluate the in-vivo effectiveness of the spinal-fixation implants with the rods made of TNTZ-ST. Following a period of 5 months, the sheep with the implanted spinal-fixation implant were euthanized and their lumbar spines

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Middle position

(A)

(B)

(C) Distal position

Fig. 4.1.5 CMRs of cross-sections of fracture models implanted with and without bone plates made of TNTZ at middle position and distal position at 48 weeks after implantation: (A) cross-section of fracture model, (B) parts of □ of (A), namely high magnification CMR of branched parts of bones formed on the outer and inner sides of tibiae, and (C) cross-sections of unimplanted tibiae.

were removed. Subsequently, undecalcified sections of the spinal region near the implanted spinal-fixation implant were ground and stained with a hematoxylin-eosin stain for histological observation. No negative effects such as screw pull-out or stress shielding of the bone could be observed in the radiographs that were captured over a period of 5 months, as shown in Fig. 4.1.6 [13]. With regard to the histological observation, no inflammatory cells could be observed in the surrounding tissue. Internal nuclei were observed in some paravertebral muscle cells. An irregular myofibular network could also be observed. In any case, no metal debris was observed but scar formation could be observed at the boundary of the spinal-fixation rods and surrounding tissue, as shown in Fig. 4.1.7 [13].

4.1.4

Improvement of the strength of low rigidity Ti alloys while keeping low rigidity

4.1.4.1 Improving static strength In the case of Ti alloys with low rigidity, the lowest rigidity, in general, obtained under solutionized conditions led to poor strength. Therefore, the strengths of such Ti alloys should be increased while keeping low rigidity. Severe cold-working processes, such

352

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(A)

(B) Sagittal plane

Coronal plane

Fig. 4.1.6 X-ray photographs of TNTZ-ST devices implanted into the spine of ovine taken from sagittal plane and coronal plane: (A) just after operation and (B) 5 months after operation.

Fig. 4.1.7 Histological image of surrounding tissue around a TNTZ-ST rod: Black region is a rod and arrows point to fibrous tissue.

100 µm

as severe cold rolling (CR) [14], severe cold swaging [15], and/or severe plastic deformation such as high-pressure torsion (HPT) [16,17], can improve the tensile strength (namely, the static strength) of β-type Ti alloys such as TNTZ. The tensile strength of the aforementioned materials can be increased to levels similar to or higher than that of Ti-64 ELI using such processes while keeping excellent ductility (elongation). This is because the high number of dislocations introduced generates a great amount of work hardening. For example [17], in the case of a TNTZ conducted with ST and subsequent severe CR (TNTZCR), the average tensile strength, 0.2% proof stress, and elongation of material were reported to be 800 MPa, 565 MPa, and 22.5%, respectively. For a TNTZ conducted with ST, the corresponding values were reported to be approximately 600 MPa, 370 MPa, and 26%, respectively. The tensile strength

Titanium spinal-fixation implants

Fig. 4.1.8 Young’s moduli of TNTZCR and TNTZHBT at rotation numbers N ¼ 1–60.

75 70 Young’s modulus, E / GPa

353

65 60 55 50 45 40

TNTZCR TNTZHPT TNTZHPT TNTZHPT TNTZHPT TNTZHPT at N = 1 at N = 5 at N = 10 at N = 20 at N = 60

and 0.2% proof stress of a TNTZ conducted with HPT (TNTZHPT) were greater than those of TNTZCR. However, the elongation of TNTZHPT exhibited a reverse trend, meaning it was lower than that of TNTZCR. Fig. 4.1.8 [17] shows the Young’s moduli (E) evaluated using a stress-strain curve obtained via tensile tests of TNTZCR and TNTZHPT at all the rotation numbers, N. After inducing severe torsional strain, the Young’s modulus of TNTZHPT slightly decreased at N < 5. That tends to be constant at N  5 as N increases. The Young’s modulus of TNTZHPT is decreased from 64 GPa to around 60 GPa for TNTZHPT at N ¼ 60, which represents a reduction of 6%. However, the Young’s modulus of TNTZHPT tends to be constant at N > 5. Solid-solution strengthening with O is also effective with regard to increasing the strength of TNTZ while keeping a low rigidity. The balance between the tensile strength and elongation of TNTZ with O contents of 0.1 mass% (TNTZ0.1), 0.2 mass% (TNTZ0.2), and 0.4 mass% (TNTZ0.4) conducted with solution treatment and aging is shown in Fig. 4.1.9 [18]. This figure also shows the tensile strength and elongation of the Ti-64 standardized in ASTM Standard F136 [19] as a reference. Although the tensile strength of TNTZ conducted with solution treatment is similar to, or lower than, that of the Ti-64 ELI alloy, the elongation of TNTZ conducted with solution treatment is greater than that of T-64. Moreover, the elongation of TNTZ conducted with aging is similar to, or lower than, that of Ti-64 but the tensile strength of TNTZ conducted with aging is greater than that of Ti-64. Furthermore, as mentioned before, the tensile strength of both TNTZ conducted with solution treatment and TNTZ conducted with aging is increased as their O content increases while their elongation is decreased. Thus, for TNTZ, the balance between the tensile strength and elongation changes remarkably due to their O content. Therefore, the tensile strength and elongation of TNTZ with a certain O content of oxygen can be controlled through heat treatment over a wide range of around 600–1400 MPa and 5%–25%, respectively, while their Young’s moduli can be kept below that of Ti-64, namely within the range of 60–100 GPa as shown in Fig. 4.1.10 [18].

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1600 Aged TNTZ0.4

1400 Aged TNTZ0.2

Tensile strength, sB / MPa

Fig. 4.1.9 Balance between tensile strength and elongation of TNTZ0.1, TNTZ0.2, and TNTZ0.4 subjected to solution treatment at 1063, 1073, and 1093 K, respectively, for 3.6 ks and aging treatment at 723 K for 259.2 ks after solution treatment.

1200 Aged TNTZ0.1

1000

Solutionized TNTZ0.4

800

Ti-6AI-4V ELI (ASTM F136)

600

Solutionized TNTZ0.2

400

Solutionized TNTZ0.1

200 0

0

5

10

15 20 25 Elongation (%)

30

35

120 Young’s modulus, E / GPa

Ti-6AI-4V ELI (ASTM F136)

100 80 60 40 20 0

Solutionized Solutionized Solutionized TNTZ0.1 TNTZ0.2 TNTZ0.4

Aged TNTZ0.1

Aged TNTZ0.2

Aged TNTZ0.4

Fig. 4.1.10 Young’s modulus of TNTZ0.1, TNTZ0.2, and TNTZ0.4 subjected to solution treatment at 1063, 1073, and 1093 K, respectively, for 3.6 ks and aging treatment at 723 K for 259.2 ks after solution treatment.

The tensile strength, 0.2% proof stress, and elongation of TNTZ with O contents of 0.14 (TNTZ-0.14ST), 0.33 (TNTZ-0.33ST), and 0.70 (TNTZ-0.70ST) mass% conducted with hot rolling followed by ST are shown in Fig. 4.1.11 [20]. As O content increases, the tensile strength and 0.2% proof stress of all TNTZ variants are increased, but there is an initial decrease in the elongation followed by subsequent increase. This trend is contradictory to those conventionally reported. The tensile strengths of the aforementioned alloys can reach values of around 1100 MPa, and elongation values of around 20% of that of TNTZ-0.70ST can be obtained. Both the tensile strength and elongation of TNTZ-0.70ST are greater than those of Ti-64

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355

28 1200

Tensile strength Elongation

24

1000 900

22

800

20

700

18

600

16

500

Elongation (%)

Tensile strength, s b/MPa 0.2% proof stress, s 0.2/MPa

26

0.2% proof stress

1100

14

400 12 300

TNTZ-0.14O

TNTZ-0.33O

TNTZ-0.70O

Fig. 4.1.11 Room temperature tensile properties of TNTZ-(0.14, 0.33, 0.70 mass%)O.

registered in ASTM Standard F136 [19]. The true stress-strain (s-s) curves of all alloys (namely, TNTZ-0.14ST, TNT-Z0.33ST, and TNTZ-0.7ST) and their work-hardening rate curves are shown in Fig. 4.1.12 [21]. When the O content increases to 0.7%, TNTZ-0.7ST shows an apparent yielding point that is not observed for TNTZ0.14ST and TNTZ-0.33ST. Furthermore, double yielding observed for TNTZ0.14ST is not observed for TNTZ-0.7ST. The work-hardening rate curve obtained for TNTZ-0.7ST showing the greatest work-hardening rate among the studied alloys intersects with the true s-s curve at true strains higher than 20%. This result indicates that this alloy shows the highest work-hardening effect among the samples, and undergoes late necking under tension. Furthermore, TNTZ-0.7ST exhibits the greatest true stress prior necking, 1300 MPa, and the greatest uniform elongation, 21%. It also exhibits an elongation to failure value of 23%, which is greater than that of TNTZ-0.33ST. Fig. 4.1.13 [21] shows scanning electron microscopy (SEM) images of TNTZ-0.33ST and TNTZ-0.7ST according to tensile tests that were interrupted at strains of 11% and 15%. Dislocation slip lines can be observed for TNTZ-0.33ST and TNTZ-0.7ST. Planar slip lines showing the same direction in individual grains can be observed for TNTZ-0.33ST, as shown in Fig. 4.1.13A. For TNTZ-0.7ST, slip lines with wavy or polyline morphologies equivalent to cross slips with multiple directions in the same grain can be observed, as shown in Fig. 4.1.13B and D. The number of slip lines that can be observed for TNTZ-0.7ST is much greater than that of TNTZ-0.33ST. The Young’s moduli of TNTZ-0.14ST, TNTZ-0.33ST, and TNTZ-0.70ST are shown in Fig. 4.1.14 [21]. The Young’s modulus of Ti-64, a conventional Ti alloy used for implants registered in ASTM Standard F136 [8,19], is also shown in this figure. The Young’s modulus of TNTZ increases as O contents increase. The Young’s moduli of TNTZ-0.14ST and TNTZ-0.33ST are lower than 65 GPa. The Young’s modulus of

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Work-hardening rate, ds /de (MPa) True stress, s (MPa)

2000 0.1ST

Work-hardening rate curves

0.3ST 0.7ST

1500

Necking point

1000

500 True stress-strain curves 0

0.00

0.05

0.10 0.15 0.20 True strain, e

0.25

0.30

Fig. 4.1.12 Room temperature true stress-strain curves and corresponding work-hardening rates of TNTZ-0.1ST (0.1ST), TNTZ-0.3ST (0.3ST), and TNTZ-0.7ST (0.7ST).

(A) 0.3ST-11%

(B) 0.7ST-11%

50 µm

(C) 0.3ST-15%

50 µm

(D) 0.7ST-15%

100 µm

100 µm

Fig. 4.1.13 (A) and (B) SEM images of TNTZ-0.3ST (0.3ST-11%) and TNTZ-0.7ST (0.7ST-11%) after the 11% strain interrupted tensile tests, respectively; (C) and (D) SEM images of TNTZ-0.3ST (0.3ST-15%) and TNTZ-0.7ST (0.7ST-15%) after 15% strain interrupted tensile tests, respectively.

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Fig. 4.1.14 Young’s moduli of TNTZ-0.14O, TNTZ-0.33O, and TNTZ-0.70O subjected to hot rolling followed by solution treatment respectively.

120 110 100 Young’s modulus, E/GPa

90

Ti-6AI-4V ELI (ASTM F136)

80 70 60 50 40 30 20 10 0

TN -0 T

0S

.7

T

3S

T

4S

.1

.3

-0

-0

TZ

TZ

TZ

TN

TN

TNTZ-0.7 is lower than 75 GPa, which is remarkably lower than that of Ti-64, namely 100–110 GPa. Therefore, for TNTZ, TNTZ with high O content exhibits high strength and elongation with low rigidity.

4.1.4.2 Improvement of dynamic strength The fatigue strength (namely, dynamic strength) cannot be improved by either severe cold working or severe plastic deformation [14]. Therefore, the introduction of a secondary phase, or secondary particles, into the β-phase matrix through aging or the direct addition of particles consisting of hard materials such as ceramics is the most effective way to improve the fatigue strength of β-type Ti alloys. It is well known that, compared with α-phase precipitation, the precipitation of the ω phase increases the strength and rigidity of the Ti alloy, but the ω phase also increases the brittleness of Ti alloys. Therefore, the fatigue strength of TNTZ is expected to be increased by introducing a small amount of the precipitation of the ω phase. For this purpose, a short aging at relatively low temperatures effectively results in a small amount of ω-phase precipitation [22]. The Young’s moduli of TNTZ conducted with ST and severe CR followed by aging at 573 K can approach values lower than 80 GPa, which is a tentative target value for a low rigidity Ti alloy for biomedical applications. The fatigue strength of TNTZ was shown to improve through an aging treatment for a period of 10.8 ks; a fatigue limit of approximately 600 MPa was

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achieved and the Young’s modulus was kept at