Performance of Metals and Ceramics in Total Hip Arthroplasty 3031254198, 9783031254192

Table of contents :
Preface
Reference
Contents
Abbreviations
1 Introduction
References
2 State of the Art in Orthopaedic Implants
References
3 Biocompatibility
References
4 Metals and Alloys Choice for Implants
4.1 Stainless Steels
4.2 Cobalt Based Alloys
4.3 Titanium and Its Alloys
4.4 Tantalum and Its Alloys
4.5 Niobium and Its Alloys
4.6 Mg Alloys
References
5 Metal Implant Allergy
References
6 Ceramics Choice for Implants
6.1 Bioactive Ceramics
6.2 Alumina Ceramics
6.3 Zirconia Ceramics
6.4 Oxide Ceramic Composites
6.5 Non-oxide Ceramics
References
7 Immuno-Allergological Compatibility Aspects of Ceramic Materials
References
8 Mechanical Aspects of Implant Materials
8.1 Metals
8.1.1 Strength
8.1.2 Wear
8.1.3 Fatigue
8.1.4 Corrosion
8.2 Ceramics
8.2.1 Hardness
8.2.2 Flexural Strength
8.2.3 Fracture Toughness
8.2.4 Fatigue
8.2.5 Wear
8.2.6 Accelerated Ageing Tests
References
9 Outlooks and Horizons in Materials and Technologies
References
10 Conclusions
Reference

Citation preview

Synthesis Lectures on Biomedical Engineering

Armando Reyes Rojas · Alfredo Aguilar Elguezabal · Alessandro Alan Porporati · Miguel Bocanegra Bernal · Hilda Esperanza Esparza Ponce

Performance of Metals and Ceramics in Total Hip Arthroplasty

Synthesis Lectures on Biomedical Engineering Series Editor John Enderle, Storr, USA

This series consists of concise books on advanced and state-of-the-art topics that span the field of biomedical engineering. Each Lecture covers the fundamental principles in a unified manner, develops underlying concepts needed for sequential material, and progresses to more advanced topics and design. The authors selected to write the Lectures are leading experts on the subject who have extensive background in theory, application, and design. The series is designed to meet the demands of the 21st century technology and the rapid advancements in the all-encompassing field of biomedical engineering.

Armando Reyes Rojas · Alfredo Aguilar Elguezabal · Alessandro Alan Porporati · Miguel Bocanegra Bernal · Hilda Esperanza Esparza Ponce

Performance of Metals and Ceramics in Total Hip Arthroplasty

Armando Reyes Rojas Laboratorio Nacional de Nanotecnologia Centro de Investigación en Materiales Avanzados S.C. Chihuahua, Mexico

Alfredo Aguilar Elguezabal Laboratorio Nacional de Nanotecnologia Centro de Investigación en Materiales Avanzados S.C. Chihuahua, Mexico

Alessandro Alan Porporati Medical Products Division CeramTec GmbH Plochingen, Germany

Miguel Bocanegra Bernal Laboratorio Nacional de Nanotecnologia Centro de Investigación en Materiales Avanzados S.C. Chihuahua, Mexico

Department of Engineering and Architecture University of Trieste Trieste, Italy Hilda Esperanza Esparza Ponce Laboratorio Nacional de Nanotecnologia Centro de Investigación en Materiales Avanzados S.C. Chihuahua, Mexico

ISSN 1930-0328 ISSN 1930-0336 (electronic) Synthesis Lectures on Biomedical Engineering ISBN 978-3-031-25419-2 ISBN 978-3-031-25420-8 (eBook) https://doi.org/10.1007/978-3-031-25420-8 © The Editor(s) (if applicable) and The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 This work is subject to copyright. All rights are solely and exclusively licensed by the Publisher, whether the whole or part of the material is concerned, specifically the rights of translation, reprinting, reuse of illustrations, recitation, broadcasting, reproduction on microfilms or in any other physical way, and transmission or information storage and retrieval, electronic adaptation, computer software, or by similar or dissimilar methodology now known or hereafter developed. The use of general descriptive names, registered names, trademarks, service marks, etc. in this publication does not imply, even in the absence of a specific statement, that such names are exempt from the relevant protective laws and regulations and therefore free for general use. The publisher, the authors, and the editors are safe to assume that the advice and information in this book are believed to be true and accurate at the date of publication. Neither the publisher nor the authors or the editors give a warranty, expressed or implied, with respect to the material contained herein or for any errors or omissions that may have been made. The publisher remains neutral with regard to jurisdictional claims in published maps and institutional affiliations. This Springer imprint is published by the registered company Springer Nature Switzerland AG The registered company address is: Gewerbestrasse 11, 6330 Cham, Switzerland

Preface

Joint replacement using different materials has been in use over the past 50 years with outstanding success. Millions of people throughout the world have benefited from total hip arthroplasty THA surgery, extending their mobility and quality of life by years. However, a clear conclusion can be the one reported by Kuncická et al. [1] pointing out that there are opportunities to improve joint replacements by improving the materials from which they are made. Materials such as metals, ceramics, and polymers serve critical functions in implant systems, and therefore their respective roles and properties are frequently shifting with the advancements in each category and the introduction of hybrid material systems. After reviewing different metallic, ceramic, polymeric, and composite materials used nowadays in THA, there remain issues that must be solved to ensure good pain relief, increasing more and more the activity levels in young patients undergoing hip replacement, as well as longevity of the prosthesis, reaching a higher range of motion as well as stability in those ranges. It was observed that the mechanical, material, and processing issues are imperative in the design, selection, and improvement in the fabrication of hip replacements. For a specific application such as THA, a correct choice of the defined material to fulfill the requirements of different standards for a particular total hip arthroplasty system component is necessary. The proper selection of metallic, ceramic, polymeric, or composite material plays a crucial role in getting the combination of properties such as high strength, wear, and corrosion resistance, as well as biocompatibility. With the passage of time and balancing the pros and cons of the various materials used for applications in THA, it could be determined that the use of ceramic materials (alumina matrix composite mainly) is best suited as bearing material for example. In conclusion, the innovations in the design and fabrication processes for the different materials for THA are raising the great reality of being able to obtain, in the medium term, implants with improved performance to match both the biocompatibility and mechanical complexity of the hip implants. However, to achieve this, surgeons in close alliance with biologists and engineers cannot be ignored. They must be persuaded of the long-term durability and reliability of the available biomaterials. Everything mentioned leads to

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Preface

the future to see novel biomaterials being developed that will increase the lifespan of orthopedic implants. Today it is not possible to fully ensure which material is the one that will dominate orthopedics. Chihuahua, Mexico Chihuahua, Mexico Plochingen, Germany/Trieste, Italy Chihuahua, Mexico Chihuahua, Mexico

Armando Reyes Rojas Alfredo Aguilar Elguezabal Alessandro Alan Porporati Miguel Bocanegra Bernal Hilda Esperanza Esparza Ponce

Reference

1. Kuncická L, Kocich R, Lowe TC. Advances in metals and alloys for joint replacement. Prog Mater Sci 2017;88:232–80. https://doi.org/10.1016/j.pmatsci.2017.04.002.

Contents

1

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

1 2

2

State of the Art in Orthopaedic Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

5 12

3

Biocompatibility . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

17 20

4

Metals and Alloys Choice for Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.1 Stainless Steels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.2 Cobalt Based Alloys . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.3 Titanium and Its Alloys . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.4 Tantalum and Its Alloys . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.5 Niobium and Its Alloys . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.6 Mg Alloys . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

23 24 26 27 29 33 34 37

5

Metal Implant Allergy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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6

Ceramics Choice for Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.1 Bioactive Ceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.2 Alumina Ceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.3 Zirconia Ceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.4 Oxide Ceramic Composites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.5 Non-oxide Ceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

59 60 62 66 69 75 78

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Immuno-Allergological Compatibility Aspects of Ceramic Materials . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Contents

8

Mechanical Aspects of Implant Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8.1 Metals . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8.1.1 Strength . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8.1.2 Wear . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8.1.3 Fatigue . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8.1.4 Corrosion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8.2 Ceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8.2.1 Hardness . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8.2.2 Flexural Strength . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8.2.3 Fracture Toughness . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8.2.4 Fatigue . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8.2.5 Wear . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8.2.6 Accelerated Ageing Tests . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

93 93 94 96 100 109 122 123 125 131 136 143 147 153

9

Outlooks and Horizons in Materials and Technologies . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

181 183

10 Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Reference . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

185 186

Abbreviations

Al2 O3 AMCs ARMD ATZ CNB CNTs CoC CoM CoP CP Ti CPTi CVD DCM DLC DMLS DNA EBSS EDS EIS ESR FDA FEA Hap HCA HCF HDP HIP HIPRBSN HIPSN HIPSRBSN

Alumina Alumina Matrix Composites Adverse Reaction to Metallic Debris Alumina-Toughened Zirconia Chevron-Notched Beam Carbon Nanotubes Ceramic-on-Ceramic Ceramic-on-Metal Ceramic-on-Polyethylene Commercially Pure Titanium Commercially Pure hcp Titanium Chemical Vapour Deposition Direct Crack Measurement Diamond-Like-Carbon Direct Metal Laser Sintering Deoxyribonucleic acid Earle’s Balanced Salt Solution X-Ray Spectroscopy Electrochemical Impedance Spectroscopy Electro-Slag Re-melting American Food and Drug Administration Finite Element Analysis Hydroxyapatite Active Hydroxycarbonate Apatite High-Cycle Fatigue Lifetime High-Density Polyethylene Hot Isostatic Pressing Hot Isostatic Pressing Reaction Bonding Silicon Nitride Hot Isostatic Pressing Silicon Nitride Hot Isostatic Pressing Sintered Reaction Bonded Silicon Nitride ix

x

HIPSSN HPSN HV ISB ISO Kf Kt LTD LTT MEM MgO Mg-PSZ MHS MoM MoP MPG MRI NaCl OCP ODH OxZr PBS PE PM PPC PVD RBSN RE SBF SCG SENB Si3 N4 SPD SRBSN SSN ST Ta THA THR TKA TZP

Abbreviations

Hot Isostatic Pressing Sintered Silicon Nitride Hot Pressing Silicon Nitride Vickers Hardness Indentation Strength in Bending International Standard Organisation Fatigue Notch Factor Stress Concentration Factor Low Temperature Degradation Lymphocyte Transformation Test Minimum Essential Medium Magnesia Partially Stabilised Zirconia Metal Hypersensitivity Metal-on-Metal Metal-on-Polyethylene Medical Product Legislation Magnetic Resonance Imaging Sodium Chloride Open-Circuit Potential Oxygen Diffusion Hardening Oxidised Zirconium Phosphate-Buffered Solutions Polyethylene Powder Metallurgy Potentiodynamic Polarisation Curves Physical Vapour Deposition Reaction Bonding Silicon Nitride Rare Earth Simulated Body Fluid Subcritical Crack Growth Single-Edge Notched Beam Silicon Nitride Severe Plastic Deformation Sintering Reaction Bonding Silicon Nitride Sintering Silicon Nitride Solution Treated Tantalum Total Hip Arthroplasty Total Hip Replacement Total Knee Arthroplasty Tetragonal Zirconia Polycrystal

Abbreviations

UHMWPE UTS YS Y-TZP ZrO2 ZTA ρ

xi

Ultra-High Molecular Weight Polyethylene Ultimate Tensile Strength Yield Strength Yttria-Stabilised Zirconia Zirconia Zirconia-Toughened Alumina Notch Radius

1

Introduction

Today, the trend in biomaterials for total hip arthroplasty is designed to tailor properties and fabricate hybrid materials to achieve excellent performance metrics in one product. Metals and ceramics are considered as biomaterials and being the central enablers of joint replacement technology in the orthopaedic field [1] without demerit the use of polymeric materials in some cases. The specific field of biomaterials has been considered and recognised as a separate entity since the first meeting held on biomaterials at Clemson University, South Carolina in 1969, and continues to receive substantial attention, exploiting the latest technologies in order to improve the quality of life and longevity of human beings [2–4]. The high performance of orthopaedic replacements is highly related with their relative motion via sliding and highly localised loads, where the wear and surface properties of both metals and ceramics are of utmost importance [5]. In reason of this, nowadays, orthopaedic implants with artificial biomaterials have allowed surgeons to ameliorate the lives of patients by reducing the pain [6] and restoring function to otherwise functionally compromised structures. The body parts most frequently replaced in the orthopaedic field are hip, knee, shoulder and spine [7], replacements for which are currently manufactured from metals like stainless steel (SS), titanium alloys, cobaltbased alloys and ceramics such as alumina (Al2 O3 ) and zirconia (ZrO2 ), or alternatively, a combination of these. However, in the case of metals, these have demonstrated strong tendencies of failure after long-term use, thus recommending the need for revision as a consequence of adverse local tissure reaction to metal debris, infection, instability, pain, aseptic loosening, and fracture, amongst others [8]. On the other hand, in spite of the advantages of the ceramic materials such as their inertness and desirable mechanical properties, such as high strength, elastic modulus, and hardness, as well as their excellent corrosion resistance, biocompatibility, low friction, high wear resistance, and stability in a physiological environment [6], ceramics have been found to be associated with some mechanical failure [9, 10]. Taking into account the advancement in medical technology, it © The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 A. Reyes Rojas et al., Performance of Metals and Ceramics in Total Hip Arthroplasty, Synthesis Lectures on Biomedical Engineering, https://doi.org/10.1007/978-3-031-25420-8_1

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1 Introduction

has become necessary to develop implants with high performance and superior longevity and biocompatibility, serving a much longer period without failure or revision surgery with the aim to satisfy the increased population of active young adults traumatised by different circumstances, so generating more comfortable living conditions. This was finally achieved with the latest generation of ceramics for THA, the alumina matrix composite with the brand name BIOLOX® delta (CeramTec GmbH, Plochingen, Germany) which was introduced in the market in 2003. Excellent reviews and original articles summarising aspects of metals and ceramics for different biomedical applications, including aspects related with orthopaedic implants, already exist. Notwithstanding the foregoing, in this review the authors emphasise a comparison between the advantages and disadvantages when both metals and ceramics are used in orthopaedic surgery, showing the evolution of these replacements provided by advances in technology and materials [11–14]. This review article, focussed on orthopaedic implants like hips and knees only, is divided into ten sections, starting with the state of the art in orthopaedic implants and their biocompatibility, the requirements to be fulfilled by metals and ceramics to be applied in the manufacture of orthopaedic implants, and the limitations and mechanical aspects of these biomaterials. The review ends looking at the horizons in materials and technologies as well as the prospects for different materials used for both temporary and permanent implants in the more used orthopaedic replacements.

References 1. Kuncická L, Kocich R, Lowe TC. Advances in metals and alloys for joint replacement. Prog Mater Sci. 2017;88:232–80. https://doi.org/10.1016/j.pmatsci.2017.04.002. 2. Geetha M, Singh AK, Asokamani R, Gogia AK. Ti based biomaterials, the ultimate choice for orthopaedic implants–a review. Prog Mater Sci. 2009;54:397–425. https://doi.org/10.1016/ j.pmatsci.2008.06.004. 3. Park JB, Bronzino JD, editors. Biomaterials: principles and applications. Boca Raton, FL: CRC Press; 2003. p. 1–241. ISBN:9781420040036 1420040030, Número OCLC, 300174064. 4. Ramakrishna S, Mayer J, Wintermantel E, Leong KW. Compos Sci Technol. 2001;61:1189– 224. https://doi.org/10.1016/S0266-3538(00)00241-4. 5. Wise DL. Biomaterials engineering and devices. Berlin: Humana Press; 2000. p. 205–319. https://link.springer.com/book/10.1007%2F978-1-59259-196-1#editorsandaffiliations. 6. Mantripragada VP, Lecka-Czernik B, Ebraheim NA, Jayasuriya AC. An overview of recent advances in designing orthopedic and craniofacial implants. J Biomed Mater Res A. 2013;101(11):3349–64. https://doi.org/10.1002/jbm.a.34605. 7. Murphy L, Helmick CG. The impact of osteoarthritis in the United States: a population-health perspective: a population-based review of the fourth most common cause of hospitalization in U.S. adults. Orthop Nurs. 2012; 31:85–91. https://doi.org/10.1097/NOR.0b013e31824fcd42. 8. Paxton EW, Fumes O, Namba RS, Inacio MCS, Fenstad AM, Havelin LI. Comparison of the Norwegian knee anthroplasty register and a United States anthroplasty registry. J Bone Joint Surg Am. 2011;93E:20–30. https://doi.org/10.2106/JBJS.K.01045.

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9. Garvie RC, Hannink RH, Pascoe RT. Ceramic steel? Nature. 1975;258:703–4. https://doi.org/ 10.1038/258703a0. 10. Rahaman MN, Yao A, Sonny Bal B, Garino JP, Ries MD. Ceramics for prosthetic hip and knee joint replacement. J Amer Ceram Soc. 2007;90(7):1965–88. https://doi.org/10.1111/j.15512916.2007.01725.x. 11. Shahadat M, Teng TT, Rafatullah M, Arshad M. Titanium-based nanocomposite materials: a review of recent advances and perspectives. Colloids Surf B Biointerfaces. 2014;126:121–37. https://doi.org/10.1016/j.colsurfb.2014.11.049. 12. Guillemot F. Recent advances in the design of titanium alloys for orthopedic applications. Expert Rev Med Devices. 2005;2:741–8. https://doi.org/10.1586/17434440.2.6.74. 13. Pezzotti G, Yamamoto K. Advances in artificial joint materials. J Mech Behav Biomed Mater. 2014;31:1–2. https://doi.org/10.1016/j.jmbbm.2013.12.012. 14. Huo MH, Martin RP, Zatorski LE, Keggi KJ. Total hip replacements using the ceramic Mittelmeier prosthesis. Clin Orthop. 1996;332:143–50. https://doi.org/10.1097/00003086-199611 000-00020.

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State of the Art in Orthopaedic Implants

The orthopaedic implants history goes back to the nineteenth century when the first ever implantation of a large human joint was performed in 1890 [1, 2], and the first total knee endoprosthesis was made of ivory and anchored to the bones using nickel-coated nails and a mixture of gypsum, resin and pumice [1, 3]. Implantation of metal-based joint prostheses began in the early 1930s with the first stainless steel total hip replacement (THR), implanted in 1938, being made of the “vanadium steel” purposely designed to be used within a human organism [4], but having inadequate biocompatibility and corrosion resistance. At present, in orthopaedic surgery, the metals more commonly used and of long tradition for joint replacements are stainless steels, cobalt-chromium alloys, titanium alloys and to a lesser extent, tantalum and its alloys, and niobium and zirconium alloys [3, 5]. These metals remain favourites owing to their good biocompatibility and their ability to undergo large plastic deformations prior to failure. Despite the above, severe limiting factors still exist for steels and Co-based alloys including fatigue of acetabular components in hip join implants leading, inevitably, to loosening of the implant [3, 6]. Indeed, with the increasing use of THR in younger and more active patients where revision rates are higher [7–9], new materials with enhanced properties or the improvement of existing ones is required. The in-situ degradation of metal-alloy implants is especially undesirable since this process may decrease the structural integrity of the implant producing an adverse biological reaction in the host leading to periprosthetic bone loss [10]. Furthermore, many authors have reported increased concentrations of local and systemic trace metals associated with metal implants [11–15]. This implies in depth studies to achieve the right balance between the implant material and its mechanical aspects, biocompatibility, high corrosion and wear resistance, and enhanced osseointegration taking into account the three main combinations of femoral head and acetabular bearing surface materials as follows: (i) metal-on-polyethylene (MoP), (ii) ceramic-on-polyethylene (CoP), (iii) ceramic-on-ceramic (CoC), highlighting that metal-on-metal (MoM) and the © The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 A. Reyes Rojas et al., Performance of Metals and Ceramics in Total Hip Arthroplasty, Synthesis Lectures on Biomedical Engineering, https://doi.org/10.1007/978-3-031-25420-8_2

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2 State of the Art in Orthopaedic Implants

combination of ceramic-on-metal (CoM) implants are uncommon [16]. The material has to guarantee the optimal performance of the orthopaedic implant without rejection to reach lower revision rates of hip arthroplasties, mainly in the younger populations, emphasising that this is one of the most common surgical procedures performed worldwide estimating, for example, that 2.5 million people are living with a hip replacement in the USA where the main indications to elect this procedure are pain and functional limitations owing to osteoarthritis [17, 18]. Additionally, in Korea, the Health Insurance Review and Assessment Service registered that more than 60,000 THAs were performed between 2010 and 2017, with an increase of more procedures over time [19]. In this context, new aspects such as bearing materials and fixation methods, along with new designs of implants, have been improved in order to return a patient to the activities of daily living and range of motion in the absence of pain [20]. Currently, the metallic biomedical materials used in orthopaedics and their limitations have resulted in improvements to the existing group of metals and the introduction in this field of ceramic materials, which will be mentioned later on. The strict use of metals in THA, that is, MoM bearings, were initially made using large head diameters during 1955–1965 [21]. Their use declined for some years in the 1970s when Sir John Charnley introduced a THA device based on MoP made-up of a small metal head and a cemented polyethylene (PE) cup [22]. Sadly, the concerns about the role of PE wear particles in osteolysis and loosening were increased with the increasing use of THA in younger and more active patients, leading to the revision rate becoming higher [23]. It is well known that stainless steel was the first class of alloy introduced for orthopaedic implants [20, 24], but independent of its ease in manufacturing and low cost, and due to its inevitable corrosion, it must only be used for short duration purposes [25]. Nevertheless, despite stainless steel being renowned as a biomaterial, it is important to stress that from the many grades in the stainless steel family, the austenitic 316L stainless steel is the only category that is used for bioimplant applications given that is favoured for its inexpensiveness, lack of ferromagnetism, exhibiting an excellent toughness, even down to cryogenic temperatures, and shows good biocompatibility according to cytotoxicity evaluation standards [26–31]. On the other hand, the mechanical properties of stainless steel can be controlled in a wide range which allows for the manufacture of products with optimal ductility and strength for different medical uses [30], guaranteeing a satisfactory mechanical performance as the material can bear significant loads and experience sufficient plastic deformation before failure, according to reports by Breme et al. [29] and Chen and Thouas [1]. However, and considering some controversial reports [32] evoking tremendous scientific interests in the fabrication of bioimplants which uses large quantities of this metal (approximately 10 to 20% in the market), in reality the mechanical working conditions inside a living body differ greatly from the external environment and stainless steel in THA is typically subjected to fatigue damage as its fatigue strength is relatively low [33–35]. In this scenario, and also taking into account its inevitable corrosion, stainless steel is now mainly being used in short-term implant devices although stainless steel

2 State of the Art in Orthopaedic Implants

7

devices remain available in other countries (particularly the United Kingdom) [20], but its applications are centred mainly in fracture plates, screws and nails, emphasising that tibial components of knee prostheses may also be manufactured from this steel [36]. Parallel with the use of stainless steel implants and owing to its superior mechanical properties and corrosion resistance when compared to other materials used at that time, the development of cobalt alloys began [1] with the first Co-based alloy implant being introduced in 1939 [4, 37]. Although the first successful implantation of a metal-on-metal hip joint replacement with this CoCr alloy was in the 1950s [38], the MoM implants tend to loosen from the bone despite continuous development and numerous technological improvements. Both 316L SS and Cr–Co alloys possess a much higher modulus than bone which in turn leads to insufficient stress transfer to the bone causing bone resorption and loosening of the implant after some years of implantation [6]. Knowing that major alloying compositions such as Co, Cr, Mo and Ni are catalogued as essential trace elements in the human body, undoubtedly they would be biologically toxic when in excessive amounts resulting in damaging to the kidneys, liver, lungs and blood cells [39–43]. Thus, the release of particles or ions caused by fretting corrosion, wear, and fatigue or loosening of the implant, is a big concern for Co-alloy biomaterials [1]. For this reason, some alloys, especially those with Ni, should be avoided in applications experiencing hard contacts and friction loading. On the other hand, the high costs involved in manufacturing have put Co-based alloys at another disadvantage when it comes to the medical market and although imperfect, cobalt alloys remain as a favourite metallic implant material in THA procedures [31]. In spite of the disadvantages, the main interest for these Co-alloys is not to design new chemical compositions, but to improve the mechanical properties and performance of the existing alloys by implementing new and novel processing technologies and thermo-mechanical treatments, as is the case of the Co–Cr–Mo ATSM F75 alloy which has been fabricated following a new technology ensuring enhanced wear resistance and fatigue strength [38]. Nowadays, Co–Cr–Mo alloys are being used along with cross-linked ultra-high molecular weight polyethylene (UHMWPE) to reduce corrosion and wear [44]. Nevertheless, fabrication and finishing treatments of these alloys have to be processed carefully in order to avoid the unintentional reduction of one property at the expense of another [38]. Independent of the above mentioned stainless steel and Co-alloys and the tendency of these metal-on-metal orthopaedic implants to loosen from the bone, despite continuous development and numerous technological improvements, Niinomi [45] reported that the natural selection of titanium-based alloys for implantation, collects a combination of outstanding characteristics such as high strength, low density (approximately 60% of the density of iron and nearly half of the density of cobalt), high immunity to corrosion, complete inertness to the body environment, enhanced biocompatibility, low modulus and a high capacity to join with bone and other tissues [45]. Thus, currently, titanium and its alloys are the most popular metallic implant biomaterials used in THA, and Ti-6Al4 V is the most commonly used alloy for stem and acetabular cementless components of

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THA [46]. Undoubtedly, titanium is superior in specific strength and has low fatigue strength, but is inferior in tribological properties compared with stainless steels and cobalt–chromium alloys, and therefore its use has increased over the past two decades because of its lower modulus, higher corrosion resistance, and better biocompatibility than both stainless steels and cobalt-based alloys [1]. However, we must not lose sight that the toxicity of vanadium has led to the development of various vanadium free Ti6Al-4 V compositions for applications such as femoral stems and fracture fixation plates where the Ti-Mo-Zr-Fe alloy is a clear example of that being practically applied for hip joint stems [1, 47]. Making a comparison of the advantages, disadvantages, and different medical applications using α, near α, α–β and β Ti alloys [48], a mechanically reliable titanium alloy orthopaedic implant has yet to become a reality. Metallic materials have predominated in orthopaedic surgery including temporary devices (e.g. bone plates, pins, and screws) and permanent implants (e.g. total joint replacements) [49], but only a few are capable of long-term success as an implant material. In fact, these metals are commonly categorised based on the major alloying element such as (i) stainless steels, (ii) cobalt-based alloys, (iii) titanium-based alloys and (iv) miscellaneous others that include NiTi and alloys of Mg and Ta [50]. Those medical implants made with this last group of metallic alloys are not yet FDA approved owing to the significant issues associated with biocompatibility [51] which is, by far, the main requirement for clinical application of any biomedical implants. Steinberg DR and Steinberg ME [52] illustrated the use of ceramics in arthroplasty during the past century and the explosive growth in the application of biomaterials in joint replacement during the past 50 years, which turned out to be the most successful surgical procedure of the twentieth century [9]. Of course ceramic materials have been integral to this evolution [53] and they were first considered for use as prostheses in the 1920s, but presenting a lack of sufficient strength and toughness to withstand the acetabular loads, and many fractured in vivo [54]. Nevertheless, and in spite of the problems, orthopaedists returned to ceramics because of favourable properties such as biocompatibility, hardness, wear, and lately, by the corrosion resistance. During this period, alumina, hydroxyapatite, zirconia, ceramic-coated metals and zirconia-toughened alumina proliferated for applications in orthopaedic implants, taking advantage of their inertness and desirable mechanical properties, as they possess high compressive strength, elastic modulus, hardness, and stability in a physiological environment [44, 55]. Ceramics used in orthopaedic surgery are classified as bioactive or inert according to the tissue response when implanted in an osseous environment. When the ceramic is inert, it merely elicits a minor fibrous reaction when it is used as bearings in total joint replacements. Conversely, when the ceramic is catalogued as bioactive, it is employed as a coating to enhance the fixation of a device, or as bone-graft substitute because of its osteoconductive properties [56]. The clinical experience has shown that ceramic bearing couples can significantly reduce particle induced osteolysis (i.e. an alumina head articulating against an alumina cup insert offers a proven solution to the problem of particle induced osteolysis by their

2 State of the Art in Orthopaedic Implants

9

very minor wear rate of under 0.005 mm per year) and the need for revision [57]. It is important to note that the most successful uses of ceramics in orthopaedics and traumatology is in the use of femoral heads and acetabular components where particularly, the hardness, and therefore the wear and scratch resistance of alumina and its composites has been clinically proven. Alternatively, zirconia ceramics were introduced in the manufacture of femoral heads for total hip replacements because of their better mechanical strength and toughness when compared to alumina [56, 57]. Approximately twenty years ago, Mg-PSZ ceramic (partially stabilised zirconia with magnesia as a stabiliser) was the very first zirconia approved by the American Food and Drug Administration (FDA). Subsequently, Mg-PSZ was replaced by Y-TZP (tetragonal polycrystalline zirconia which is stabilised by yttria) owing to its better mechanical strength. However, problems in the clinical use of zirconia heads related to its lack of phase stability under hydrothermal conditions [58, 59] have led to high revision rates [60]. It is worth remarking that friction in the hip joints results in a not insignificant temperature increase. Nevertheless, Lu and McKellop [61] reported an increase in the temperature of the wear couple zirconia/polyethylene compared with that of alumina/polyethylene due to the poor thermal conductivity of zirconia. Considering this inconvenience, the failure risk of zirconia is high and as such the wear couple zirconia/zirconia has not been introduced in THA [53, 62–64]. After the zirconia failure, the development of ceramics for THA moved to alumina matrix ceramic composites as the zirconia-toughened alumina (ZTA) involving the combination of the attractive properties of both alumina and zirconia where alumina offers hardness, wear resistance, and phase stability while zirconia offers attractive fracture toughness. As a result of this, several research projects aimed to develop a suitable ZTA material for THA, such as Toni’s team [65, 66]. The first ZTA for THA was introduced by CeramTec with the brand name BIOLOX®delta, which is today considered the gold standard bearing material for THA. Even if the market is led by a ZTA material produced by a single manufacturer, it is worth mentioning that there are diverse manufacturers for femoral heads and cup inserts for use in THA, as well as package leaflets and instructions that contain information about which implants and combinations are approved. Therefore, if an application or combination of components have been permitted, then research reports on the functional efficiency on which the decision to release the combination was approved, will exist. Based on the latter, it is of paramount importance keep in mind and according to a report by Willmann [57] that if a surgeon combines components which are not explicitly approved, then he becomes fully accountable, in accordance with medical product legislation (MPG), for quality assurance. Moreover, the surgeon is neither trained nor does he have the equipment to prove his case considering that there are no technical investigations in existence which could serve as the basis for approval certification. On the other hand, it is of paramount importance keeping in mind that combinations of components from different manufacturers are not approved. Therefore, the recommendation is never to be mixed and matched.

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Table 2.1 Candidate metallic materials for femoral components and their compositions. From Ref. [71]. Bahraminasab M, Bin Sahari B. Published in “Shape Memory Alloys, Processing, Characterisation and Applications” under CC BY 3.0 license © 2013 Bahraminasab and Bin Sahari, licensee InTech. This is an open access chapter distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited Material number Material’s names

Compositions Fe balancing, 17–20% Cr,10–14%Ni, 2–4% Mo, 0.03–0.08% C, 2% Mn and 0.75% Si

1

Stainless steel L316 (annealed)

2

Stainless steel L316 (cold worked)

3

Co–Cr alloys (Wrought Co–Ni–Cr–Mo) Co balancing, 19–21% Cr, 9–11% Ni, 14.6–16%W, 0.13% Mo, 0.05–0.15%C, 0.48%Si, and maximum 2%Mn &3%Fe

4

Co–Cr alloys (Castable Co–Cr–Mo)

Co balancing, 27–30% Cr, 2.5% Ni, 5–7% Mo, 0.75%Fe, 0.36%C and maximum 1%Mn &Si

5

Ti alloys (Pure Ti)

0.3% Fe, 0.08% C, 0.13% O2 , 0.07% N2

6

Ti alloys (Ti-6 Al-4 V)

Ti balancing, 5.5–6.5% Al, 3.5–4.5% V, 0.25% Fe and 0.08% C Ti balancing, 5.50–6.50% Al, < =0.080% C, 500 MPa and a UTS ranging between 900 and 1540 MPa, values comparable to the properties of the new generation of steels. Here again the manufacturing process has a great influence on the resistance of CoCr alloys since the strength is generally higher for wrought than for cast work pieces [4]. The YS for cast Co–Cr–Mo used in biomedical applications (ASTM F75) ranges from 450 to 520 MPa, while the UTS lies between 650 and 890 MPa. On the other hand, forged Co–Cr–Mo (ASTM F799) has YS levels between 900 and 1030 MPa and UTS levels from 1400 to 1590 MPa. Higher values can be obtained after thermomechanical processing in the other Co–Cr–Mo types alloyed with elements such as W and Ni (e.g. ASTM F90 and F562). Likewise, the highest YS value amongst all the materials used for THA and TKA, has been reached by means of the cold processed and aged ASTM F562 Co–Ni–Cr–Mo alloys (almost 1800 MPa) [11]. However, and in spite of these high YS values, the highest specific strength amongst all the metal biomaterials used in orthopaedic implants is observed in Ti-based alloys. It is important to point out that titanium has been a major game changer in the prosthetics industry [12]. Based on its benefits as well as disadvantages, titanium is a significant metal in the orthopaedic field that is worth building upon and improving. Aware of the

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shortage of long-lasting implants, engineering is constantly working with the purpose to create new alloys of certain metals that benefit the patient [13]. There are numerous titanium alloy types used in orthopaedic applications. In this context, the mechanical properties and strength will depend, in great measure, on the chemical composition and phase structure [9]. In general, these mechanical properties are comparable to stainless steel, but they are lower than for Co-based alloys. Depending on the impurity content, the YS value varies between 170 and 480 MPa, whereas its UTS varies from 240 to 550 MPa [4]. In the case of Ti–6Al–4 V, its UTS can reach 930 MPa for similar vanadium-free alloys, while the YS are typically between 795 and 860 MPa [4]. On the other hand, β phase alloys commonly have a UTS from 600 to 1100 MPa with the YS achieving similar maximum values. Niinomi et al. [9] reported that the strength of Ti alloys increases with increasing oxygen content and higher values can be obtained following various treatments, such as solution treatment with subsequent ageing, thermo-mechanical processing and severe plastic deformation (SPD) technologies [11]. Typical examples of the optimisation of mechanical properties by thermo-mechanical processing followed by solution treatment and ageing process are the two most important β Ti alloys used nowadays in orthopaedic implants, Ti–29Nb–13Ta–4.6Zr (wt. %) alloy and the Ti–13Nb–13Zr (wt. %) alloy. It is of paramount importance to stress that despite their advantages and promise, these β Ti alloys have only found use in niche applications since their extraction, processing, and alloying are directly associated with a high cost according to studies by Kolli and Devaraj [14]. However, there are still controversies regarding which is the best titanium alloy for orthopaedic replacements. Regardless of the two Ti alloys mentioned above, the Ti–6Al– 4 V alloy is an optimal choice for THA due to its excellent strength, flexibility, fixation as well as osseointegration, in addition to its lightweight and adaptation to different body temperatures and room temperatures [15]. However, although the range of application of Ti-based alloys in the medical field and orthodontic surgery is truly astonishing (joint replacement parts for THA, TKA, shoulder, spine, elbow, to name a few [2, 16], in general terms, the requirements of strength in these alloys for biomedical applications are much lower and have received less attention in the literature. Instead, they receive more attention regarding properties such as modulus of elasticity, cytotoxicity, corrosion resistance, and biocompatibility [14, 17, 18]. Table 8.1 [2, 4, 19–24] summarises the YS (stress which will cause a permanent deformation of 0.2% of the original dimension) and UTS (maximum stress a material can withstand) values for the main titanium and titanium alloys used in orthopaedic implants. As seen from Table 8.1, CP–Ti has the YS and UTS in the range of 275–480 MPa and 345–550 MPa, respectively [25]. However, owing to low mechanical strength at room temperature, α-type Ti alloys have not been used as implants [4] although they belong to the first generation of CP–Ti developed between 1950 and 1990, and nowadays are widely used as biomaterials for dental and medical applications since they can promote the rapid osteointegration [2]. On the other hand, these α-type Ti alloys are prohibited for use in load-bearing applications as a consequence of their low mechanical strength and

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Table 8.1 Yield strength (YS) and ultimate tensile strength (UTS) values for the main titanium alloys used in orthopaedic implants [2, 4, 19–25] σ0.2 (MPa)

Materials

σUTS (MPa)

Type of alloy

CP–Ti grade 2

275

345

α

CP–Ti grade 3

380

445

α

CP–Ti grade 4

480

550

α

Ti–6Al–4 V (annealed)

825–869

895–930

α+β

Ti–6Al–4 V ELI (mill annealed)

795–875

960–965

α+β

Ti–5Al–2.5Fe

820

900

α+β

Ti–15Mo–5Zr–3Al (aged)

771

812

β

Ti–12Mo–6Zr–2Fe (annealed)

1000–1060

1060–1100

β

Ti–36Nb–2Ta–3Zr–0.3O

670–1150

835–1180

β

Ti–6Mn–4Mo

1090

1105

β

Ti–35Nb–5Ta–7Zr

530

590

β

σ0.2 and σUTS are yield strength (YS) and ultimate tensile strength (UTS)

low fatigue strength [26]. Among the α+β-type Ti alloys, the Ti–6Al–4 V (Ti64) is the most widely and commonly used, accounting for 50% of the total Ti production besides CP–Ti [27] This alloy in the annealed condition is the most used for orthopaedic and traumatology implants [28]. Referencing the β-type Ti alloys [23], these possess comparable strength and better biocompatibility and therefore, are expected to exhibit higher corrosion resistance in the human body compared with (α+β)-type Ti [29]. Since these β-types Ti alloys are alloyed with elements such as Mo, Ta, Nb, and Zr, their density is high. These alloying elements are rare which leads to a high cost of raw materials. Moreover, the high melting points required for manufacture cause segregation and difficulty in preparation of these alloys [23, 30]. In light of this, elements such as Fe, Mn, Sn, and Cr are being used in order to develop new low-cost β-type Ti alloys with favourable mechanical properties for biomedical applications [31, 32].

8.1.2

Wear

Over the years, different metals have been used in a variety of applications in the medical field and more specifically in orthopaedic implants [33]. Amongst the metals and alloys primarily used in biomedical applications are stainless steels, Co alloys, and Ti alloys [34, 35]. It is well known that wear and corrosion are the primary reasons for the failure of implant elements [36]. Regardless of some applications of tribology in the biomedical field such as wear of dentures [37], heart valves [38], plates as well as screws in bone fracture repair [39], to name a few examples, it is without doubt that the wear is the most

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important factor in controlling and determining the long-term clinical performance of a metallic biomaterial as a replacement which can be influenced by factors such as material properties, conditions of the type of alloys developed, fabrication processes, experimental test methods, and characterisation methodologies, amongst others [36]. According to Prasad et al. [40], the degradation of implants is initiated from mechanical wear of the components as well as through electrochemical corrosive reactions with body fluids, with mechanical wear being the most prominent in MoM implants, and particularly in THA. Factors related to the patient’s life including its quality and quantity of synovial fluid, level and type of stresses on articulating surfaces, material properties, geometry and dimension, as well as the employed surgical technique, are of paramount importance to take into account to evaluate the wear of the components in contact with artificial joints [41, 42]. The abrasive, adhesive, fatigue, and corrosive modes are considered the four types of mechanism that may be implicated in the wear of metal components of joint prostheses [43]. It is expected that adhesive wear would be the most prominent in the MoM combinations, whereas abrasive wear is catalogued as the predominant mechanism in the all cobalt-chromium alloys where the socket of these prostheses also showed evidence of considerable adhesive wear, while fatigue wear was noted on one component according to the study by Walker et al. [44]. The prostheses combinations can be MoM, CoC or MoP. In these combinations the primary reasons for material related revision surgeries are aseptic loosening, implant breakage, and wear of the acetabular, which amount to 58% [45]. Today, mainly MoP, CoP and CoC are used whilst MoM is banned. Wear in metallic orthopaedic joints occurs in four different modes. An excellent approach to the wear modes is described by Seyyed Hosseinzadeh et al. [46]: “Mode 1 wear results from the motion of two primary bearing surfaces one against each other, as intended. Mode 2 refers to the condition of a primary bearing surface moving against a secondary surface, which is not intended. Usually, this mode of wear occurs after excessive wear in mode 1. Mode 3 refers to the condition of the primary surfaces moving against each other, but with third body particles interposed. In mode 3, the contaminant particles directly abrade one or both of the primary bearing surfaces. This is known as three-body abrasion or three-body wear. The primary bearing surfaces may be transiently or permanently roughened by this interaction, leading to a higher mode 1 wear rate. Mode 4 wear refers to two secondary (nonprimary) surfaces rubbing together”. The Fig. 8.1 clearly illustrates these wear modes. Keeping in mind that the wear resistance of a metallic replacement is clinically very important, several studies have been performed in order to investigate their tribological properties. For this, different wear test methods have been used by various researchers to characterise metallic prostheses [36]. However, the characterisation of the wear and friction of different metallic biomaterials using a suitable test methodology is critical. Reviewing the scientific literature, the methods most commonly used to study the tribological behaviour of metallic biomaterials are the block-on-disc [47, 48], ball-on-disc

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8 Mechanical Aspects of Implant Materials

Fig. 8.1 The modes of wear for a total hip arthroplasty: A Mode 1 or normal wear, B Mode 2 or subluxation wear, C Mode 3 or third body abrasive wear, D and E Mode 4. From Ref. [46]. Seyyed Hosseinzadeh HR, Eajazi A, Shahi AS. The Bearing Surfaces in Total Hip Arthroplasty—Options, Material Characteristics and Selection, Recent Advances in Arthroplasty, Dr. Samo Fokter (Ed.), ISBN: 978-953-307-990-5, InTech, 2012. Under the terms of the Creative Commons Attribution 3.0 Licence. © 2012 The Author(s)

[49–52], and pin-on-disc [53] where such tests are commonly conducted in an environment of simulated body fluids (Ringer’s solution) [47], or fluids containing NaCl and phosphate-buffered solutions (PBS) [54]. Table 8.2 [36, 55] summarises the advantages and disadvantages of various wear test configurations. The great variety of investigations concerning the wear of biomedical implants have demonstrated that although the most common type of hip joint comprises femoral head articulating against an UHMWPE acetabular cup, MoM prostheses produce 20–100 times lower wear volumes compared to MoP bearings [56]. For example, McGee et al. [57] noted that titanium alloy femoral heads consistently had the maximum wear, averaging 74.3% against UHMWPE acetabular components. In this context there is a resurgence of the MoM bearings instead of MoP, however, these produce nanometre-to submicrometric sized metal particles [58, 59] which lead to a large cumulative surface area for corrosion, inducing pathological changes that are associated with hypersensitivity and osteolysis [60, 61]. With the purpose to overcome this wear related and hence, the revision surgery, there has been continuous effort to replace the cup material from polymer to metal or ceramic, thus, the long-term problems associated with UHMWPE wear debris have led to the exploration of the use of MoM prostheses. It was observed [2, 62] that the biological

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Table 8.2 Advantages and disadvantages of various wear test configurations [55]. From Ref. [36]. Hussein MA, Mohammed AS, Al-Aqeeli N. Wear characteristics of metallic biomaterials: a review. Materials 2015;8:2749–68. Licensee MDPI, Basel, Switzerland under the terms and conditions of the Creative Commons Attribution licence (http://creativecommons.org/licenses/by/4.0/) © 2015 by the authors Test

Advantages

Disadvantages

Pin-on-disc

After run-in, surface pressure remains constant Easy to determine wear volume and wear rate The model closely simulates a linear friction bearing

Difficult to align pin. If the pin does not stand perfectly vertical on the plate, edge contact results A very long run-in time is therefore necessary The front edge of the pin can skim off lubricant This makes a defined lubrication state impossible

Ball-on-disc

High surface pressures are possible The ball skims off lubricant less than a pin does The model is similar to a linear friction bearing and a radial friction bearing

Very small contact ratio: The contact surface of the ball is small compared to the sliding track on the disc The contact area is enlarged by wear Difficult to determine the wear volume of the ball

Block-on-disc The model is capable of simulating a variety of harsh field conditions, e.g., high temperature, high speed, and high loading pressure

reaction to metal particles in vivo is markedly different to that produced by UHMWPE wear debris, and lower inflammatory reactions are caused by metal [63]. In spite of MoM prostheses exhibiting higher frictional torques than the MoP [64], the problem remains latent for the effect of the metal particles released after long durations [2]. The optimum material combinations to give the lowest wear rates coupled with the greatest host tolerance to the wear particles remains a challenge. Considering the options MoP or MoM, currently there is not definite evidence to favour one or the other, although MoM has been proven to be a disaster. However, there is evidence that cobalt-chromium particles may be more harmful than those of polyethylene [43, 65, 66]. Moreover, neither stainless steel against itself nor titanium against itself should provide a low-wear rate, lowfriction system because of their tendency to gall and seize [43]. Inevitably, wear debris generates at all sliding interfaces owing to the relative motion of the implanted components (between polymers, ceramics, and metals) inasmuch as all the materials have a certain finite roughness [11]. What is a reality is that wear for MoM and metal-on-ceramic (MoC) designs is far lower than at interfaces with polymers; therefore, the largest amount of debris is generated by polymers sliding against metals and ceramics. The CoC designs exhibit a volumetric wear rate of approximately 0.04 mm3 /year, while the rate for the best

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MoM design was 0.9 mm3 /year [67]. Wear rates for polymer-metal hip implant systems were reported to be between 4.5 mm3 /year and 78 mm3 /year [67–69]. It is important to stress that in vitro, CoM has also been banned. Considering that MoM are the bearing couples more used in THA, [59, 60, 62, 70–73], wear in MoM bearings can result in scratching and eventually erosion of the material, thus, wear and corrosion are very probably the major causes of release of metal into the tissues of MoM patients. Based on the mean time in situ, an annual wear rate of 7.6 mm/year (range, 2.9–13 mm/year) was calculated for the bearing combinations. Likewise, also based on the implants’ time in situ, the mean volumetric wear rate was 2.02 mm3 /year (range, 0.55–3.74 mm3 /year). Now, with a density of the alloy of 8.38 g/cm3 , the mean gravimetric wear rate was calculated to 17 mg/year (range, 4.6–31 mg/year) [70]. On the other hand, it is believed that the linear wear rates are dependent on a multitude of factors such as the type of implant and positioning, according to reports by Onda et al. [74], Shimmin et al. [75], and Williams et al. [76]. Table 8.3 presents an excellent collection of various developed materials and their manufacturing techniques, the types of wear tests carried out, as well as the parameters used to characterise wear, along with the results obtained in each investigation [47–53, 77–85]. There are interesting techniques to improve the tribological properties of metal replacements, altering the nature of its surface by applying modern concepts of surface engineering [41] such as ion implantation (physical deposition) to modify the physical and/or chemical properties of the surface (this technique has been reported to improve the wear performance of Ti–6Al–4 V and Co– 28Cr–6Mo alloys) [86], nitriding (a thermo-chemical surface treatment) used to increase the resistance of a Ti–6Al–4 V alloy to dry wear [87, 88], carburisation and boriding techniques used to enhance surface hardness, which in turn improves the wear resistance of different metallic alloys, and plasma spray coating to enhance the wear resistance of a few biomaterials, to mention some [36].

8.1.3

Fatigue

In modern medicine, the use of metallic implants is more and more important [89]. Fatigue fracture as well as are some of the major problems associated with implant loosening, stress-shielding, and ultimate implant failure [90, 91]. It is well known that under diverse physiological conditions, orthopaedic implants are subjected to a very complex interaction of mechanical and chemical–biological loading components [92] where the cyclic loads, which have to be sustained in a very corrosive medium containing enzymes, proteins and cells, are one aspect of the in vivo loading scheme. Generally, fatigue takes place when repetitive cyclic loading on the implant weakens the surface leading to the production of cracks, irremediably causing fragmentation and pitting, including in some cases, gouges, etches, surface discolouration, surface deposits and third-body particulate generation [40, 93]. Additionally, Teoh [90] and Geetha et al. [2] reported the occurrence of fatigue failure

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Table 8.3 Wear studies of metallic biomaterials. From Ref. [36]. Hussein MA, Mohammed AS, Al-Aqeeli N. Wear characteristics of metallic biomaterials: a review. Materials 2015;8:2749–68. Licensee MDPI, Basel, Switzerland under the terms and conditions of the Creative Commons Attribution licence © 2015 by the authors First author, references

Material & fabrication processes

Experimental test techniques & parameters

Main results

Cvijovic et al. [47]

Ti–13Nb–13Zr Ti–6Al–4 V Arc melting

A block-on-disc tribometer was used to conduct wear and friction tests in a simulated body fluid (Ringer’s solution) Temperature: ambient Normal load: 20–60 N Sliding speed: 0.26–1.0 m/s

The Ti–6Al–4 V alloy showed a higher wear resistance than the Ti–13Nb–13Zr alloy Abrasion was the primary wear mechanism

Gialanella et al. [48]

NiTi Commercial alloy

A block-on-disc was used to measure dry sliding wear. A profilometer was used to quantify wear Sliding speed: 0.837 ms−1 Sliding distance: 1004 m Loads: 50 to 200 N

A NiTi/WC–Co coupling exhibited a high wear rate Wear mechanism: a transition from delamination wear to a regime featuring a mixture of delamination and oxidation wear

Suresh et al. [49]

Ti–13Nb–13Zr Equal channel angular pressing (ECAE)

A tribometer was used as a lubricity fretting test system for texture and wear behaviour; fretting wear and 3D surface texture measurements were performed Normal loads: 6 N Frequency: 20 Hz Temperature: 37 ± 0.1 °C

The grain size and the texture of material affected the wear of the surface There was no difference in the friction coefficient between the ECAE processed and as-received samples

Xu et al. [50]

β-type Ti–15Mo–xNb arc-melting vacuum-pressure casting system

A ball-on-disc was used for dry wear tests Normal load: 1 N and 2 N Test-disc rate: 100 r/min

The lowest friction coefficient was obtained for a Ti–15Mo–5Nb alloy under a 1-N load Adhesion was the primary wear mechanism (continued)

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Table 8.3 (continued) First author, references

Material & fabrication processes

Experimental test techniques & parameters

Main results

Fellah et al. [52]

Ti–6Al–7Nb and AISI 316L stainless steel

Ball-on–disc and sphere-on-plane Load: 3 N, 6 N and 10 N Sliding speed: 1, 15 and 25 mm/s

The same mechanisms of wear and friction were found for all of the tested samples

Li et al. [53]

Ti–Nb–Ta–Zr and Reciprocal pin-on-disc in Ti–6Al–4 V induction a 0.9% NaCl solution skull melting method Reciprocating velocity: 45 rpm Sliding distance: 30 km

The wear resistance of Ti–29Nb–13Ta–4.6Zr was enhanced by incorporating Nb2 O5 oxide particles into the diffusion-hardened surface of the alloy

Choubey et al. CP Titanium, [77] Ti–6Al–4 V, Ti–5Al–2.5Fe, Ti–13Nb–13Zr and Co–28Cr–6Mo

Ball on flat fretting wear tester: Hanks’ balanced salt solution Normal load: 10 N for 10,000 cycles Frequency: 10 Hz

The primary wear mechanisms of Ti alloys were tribomechanical abrasion, transfer layer formation and cracking

Iwabuchi et al. Co–29Cr–6Mo alloy [78] and Ti–6Al–4 V

Fretting apparatus and a reciprocating sliding tribometer: Quasi-body fluid, Hanks’s solution Normal load: 5 N; frequency: 10 Hz Temperature in the solutions: 37 ± 2 °C

Co alloy exhibited good wear resistance; Ti alloy exhibited good fretting resistance

Luo et al. [79] ASTM F1537 Co–Cr alloy

Pin-on-disc tribometer Load: 20 N Rotation speed: 60 (rpm)

The tribo-corrosion properties of the Co–Cr alloy were enhanced by a layer of the S-phase

Chiba [80]

Co–Cr–Mo forged

Pin-on-disc Load: 9.8 N 24 rpm

Forged CoCr exhibited a lower wear loss than a cast CoCr alloy

Chan et al. [81]

(CoCr), stainless steel Pin-on-disc Sliding speed: 0.5 mm/s, (SS) 5 mm wear track radius Normal load: 1.8 N (continued)

8.1

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103

Table 8.3 (continued) First author, references

Material & fabrication processes

Experimental test techniques & parameters

Main results

Fischer et al. [82]

AISI 316L Co–Cr29–Mo6

Pin-on-disc for dry sliding wear tests Load: 5 N Relative velocity: 0.1 m/s Ambient temperature: 25 °C

Ni-free high-nitrogen steel and LC–CoCrMo alloy exhibited higher wear resistance and dry friction than Ni-containing austenitic steels

Muñoz [51]

Co–Cr–Mo Tribo-corrosion Low and high carbon techniques Load: 1.2 N; frequency: 1 Hz Temperature: 37 ± 0.1 °C Simulated body fluids [NaCl and phosphate-buffered solutions (PBS) with and without albumin]

LC CoCrMo had a higher wear resistance in NaCl and PBS albumin than HC. No differences were observed for the alloys in the other solutions

Alvarez-Vera et al. [83]

Co–Cr alloy with boron additions (0, 0.3, 0.6 and 1 B wt.%) by casting method

Three-axial hip joint simulator

Wear resistance as the boron increased

Mohan and Anandan [84]

Commercial Ti–13Nb–3Zr alloy oxygen implanted

Reciprocating type wear tester normal forces: 3, 5 and 10 N The stroke length: 10 mm and an alumina ball of 6 mm diameter was used as the counter surface

The implanted samples display a lower friction coefficient as compared to the substrate one

Attar et al. [85]

Commercially pure titanium (CP-Ti) parts produced using selective laser melting (SLM) and casting

A pin-on-disc at room temperature A stainless-steel disc of 45 mm diameter Loads: (15 N, 20 N,25 N and 30 N) Sliding speed: 0.5 m/s for 15 min

SLM CP-Ti showed better wear resistance compared to casting as a result of fine grains and higher microhardness

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of hip implants when they are subjected to cycles of loading and unloading over many years. In this context, it is likely that the majority of mechanical failures occurring in orthopaedic implants during use are fatigue failures, and that it is equally probable that the saline environment influences this fatigue behaviour. For the case of metallic joint replacements, this has to transmit load unaided by bone and so the occurrence of fatigue failures is far more serious [43]. However, and even after several years of widespread clinical use in THA, few reports related to mechanical failures have been reported because the conditions under which these failures occur are not clear. Nevertheless, in spite of reports by Ducheyne et al. [94] describing eight cases where everyone involved a prosthesis in a valgus position, a more in depth study is required in order to determine an approach conducive to understanding the mechanisms that cause the catastrophic failure by fatigue of the metallic prosthesis. In general, failures have been reported in stainless steel, cobalt-chromium alloys, as well as titanium and titanium alloys. It is worth noting that together with wear, fatigue is among the most harmful phenomena which can result in the ultimate failure of orthopaedic implants [11, 90]. The fatigue strength, which is related with the maximal cyclic stress before the occurrence of any damage or failure, depends principally on the microstructure, surface quality of the component, and the loading conditions [4, 90]. It usually initiates in the most highly stressed location, and therefore, is especially dangerous since it can occur before the stress level reaches the UTS and even the YS, without any warning. It is important to differentiate between fretting fatigue (hip joints mainly) that occurs when the cyclic stress combines with the influence of friction, and fatigue that is supported by the effects of corrosion and commonly referred as corrosion fatigue [4]. Recent premature failures of metal THA have raised concerns in the medical community about their load bearing capacity, safety, reliability, and survival rates, inasmuch as specifically, the femoral stem fracture is one of the most acute complications resulting in greater morbidity and higher cost of revision hip surgery. However, Nganbe et al. [95] reported an unusual case of fatigue failure of a cobalt-chromium alloy cementless femoral stem after 24 years of excellent performance, where the failure occurred at the neck-stem junction remote from any modular interface, and the fractographic studies and microstructural investigations revealed no intrinsic defect of the failed component. This behaviour illustrates the importance of proper long-term testing of implants in regard to cyclic loading since today the patients are younger and more active, and therefore, these implants must guarantee a long-term fatigue resistance. Despite numerous investigations on the fatigue strength of common metallic implant alloys used in orthopaedic applications, such as stainless steel, cobalt, and titanium alloys, the occurrence of failures caused by fatigue is still a major problem [96– 98]. Microstructural characterisations have shown a broad spectrum of characteristics on fracture surfaces of failed implants, such as intergranular/transgranular cracking, fatigue striations, and cleavage or cleavage like surface topography. In the scientific literature there are numerous results of fatigue strength in metals commonly used in orthopaedic

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replacements which can vary depending on the technique used to characterise such fatigue strength [86, 89, 95, 96]. In the field of metals used for implants, for example, the fatigue strength of as-cast stainless steels in aqueous biological solutions is usually between 200 and 300 MPa [4, 11], this being acceptable for most applications within the human body. Nevertheless, these values can be improved via subsequent deformation processing and/or by nitrogen content increase. Likewise, the fatigue strength of forged 316L steel (steel most commonly used metallic material in orthopaedic implants) was reported to reach 400 MPa according to investigations by Teoh [90]. Nonetheless, Co-based alloys generally have higher fatigue strengths than stainless steels, and taking into account this advantage, CoCr alloys are a suitable choice for applications for cyclically loaded prostheses and any highly loaded component reaching fatigue strengths > 500 MPa, or alternatively, 700 MPa reached by means of modern powder metallurgy (PM) techniques such as Hot Isostatic Pressing (HIP) [90]. However, the processing route influences the fatigue strength of these Co-based alloys, obtaining a value as low as 100 MPa for cast and 200 MPa for mechanically processed components. Continuous investigations have led to improvements in casting methods, inspection, and machining techniques to significantly increase the fatigue strength of Co-based alloys. Thanks to these improvements, implants manufactured from the Co-based alloys now enjoy a wide acceptance throughout the medical world since a great deal of data has been accumulated from both research and clinical studies that justifies much of the confidence in applying these type of alloys in orthopaedics [99]. However, an important observation must be taken into account, which is the fact that these implants experience a severe cyclic loading, negatively affecting their fatigue performance for the specific case of cast and forged Co alloys under simulated body fluids, which can be the cause of the low success of THA after 20 years [4, 100]. Indeed, the fatigue resistance of forged Co-based alloys is considerably higher than that of cast alloys, the fatigue strength of forged Co–Cr being > 600 MPa (in air), whereas for cast Co–Cr, it is approximately 300 MPa [4, 90, 101]. The fatigue behaviour of Co-based alloys for orthopaedic implants can be improved by controlling the alloy composition, and via heat treatment [102]. For example, Hernandez-Rodriguez et al. [100] increased the fatigue resistance in the as-cast condition of Co-based alloys via the addition of boron (B) and they observed a fatigue resistance of 615 MPa for the alloy with 1 wt. % B, representing an improvement of 60% regarding the base alloy. Conversely, heat treatment diminished the fatigue resistance of most alloys containing B, except the 0.06 wt. % B, which retained its fatigue resistance at the same level of 433 MPa, as observed in the as-cast condition. Moreover, continuous investigations have demonstrated that by means of surface treatments [11], the fatigue strength and other properties of the Co-based alloys can be significantly improved, as will be described later on. Concerning titanium alloys, the Ti–6Al–4 V (Ti64) alloy is still the most commonly used α+β titanium biomedical alloy and is used under the annealed condition whereas the metastable biomedical alloys are preferred in the solution treated (ST) and, ST and aged

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conditions [2]. It is important to remember the urgent need to find more reliable materials, metals, and mainly ceramics, to replace broken or deteriorating parts of the human body due to the rapid increase in the number of both younger and older recipients. In this context, there is a necessity to obtain metals and alloys of extreme chemical inertness and adequate mechanical strength [103, 104]. Therefore, the medical community has found an opportunity with the use of titanium and its alloys in orthopaedic replacements such as THA and TKA, taking advantage of their excellent corrosion resistance, a high strength to weight ratio, and low elastic modulus [105]. Regardless of the material used in orthopaedic implants design, the presence of geometrical fillets such as notches, which cause local stress, is unavoidable and therefore, it is necessary to pay attention to these notches, considering that they affect the fatigue resistance [106]. Due to this, the failure frequently takes place as a result of high tensile stresses at the surface or around notches (for example drill holes in intramedullary nails or plates) [107]. Another important point to take into account is the interdependency between factors such as implant shape, material, processing technique, as well as the type of cyclic loading, making the determination of the fatigue resistance of a component an intricate, but critical task, according to the report by Hosseini [103]. The titanium alloy Ti–6Al–4 V is generally considered as a standard material when evaluating the fatigue resistance of orthopaedic titanium alloys. However, its mechanical response is extremely sensitive to its prior thermo-mechanical processing history (prior β grain size, the ratio of primary α to transformed β, the α grain size, the α/β morphologies, surface preparation) which especially impacts the high-cycle fatigue lifetime (HCF) [103]. In fact, enhancement of the HCF resistance of Ti–6Al–4 V may be achieved, under careful control, by shot peening [103] where, under a cold working process, the surface is bombarded with typically spherical media so plastically deforming the surface and then the resulting compressive residual stresses may provide an increased part life when the surface related failure mechanisms, such as fatigue or corrosion, are involved [108]. However, to guarantee the optimum performance, a perfect balance between, on the one hand, the high compressive surface residual stresses and on the other hand, the increased surface roughness produced during shot peening, must be achieved. This surface treatment increases the fatigue strength of Ti–6Al–4 V by 10% over grit blasting alone. The effect of different surface finishing techniques on the fatigue strength of both α and α/β titanium alloys is illustrated in Table 8.4 [103, 109]. According to Dowling [110], the denominated fatigue notch factor Kf, (Table 8.4) is affected by the notch radius (ρ). Specifically, when Kf = 1, there is no degradation in the component due to the notch, whereas if the notch has a large radius (ρ) at its tip, then Kf may be essentially equal to Kt (stress concentration factor which depends upon the geometry of the notch and the loading mode, and is defined by the maximum local stress (σmax ) to nominal stress (σnom ) ratio) [111] and in this scenario the notch produces the full theoretical reduction in the fatigue limit. However, it is important to highlight that owing to the complexity involved,

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Table 8.4 Effect of surface preparation on the fatigue of α and α/β titanium alloys [109]. From Ref. [103]. Hosseini S. Fatigue of Ti–6Al–4 V, Chapter 3 in Biomedical Engineering—Technical Applications in Medicine, Radovan Hudak, Marek Penhaker and Jaroslav Majernik, IntechOpen, 2012, pp. 75–92. Under the terms of the Creative Commons Attribution Licence © 2012 Hosseini. Licensee InTech Titanium alloy

Test conditions

Fatigue limit (MPa)

Kf a

cpTi

Mechanically polished

234

–

Ti–6A1-4 V

Electrolytically polished

200

0.9

Gentle surface grinding

427

–

Gentle chemical machining

90

0.8

Abusive chemical machining

352

0.8

Abusive surface grinding

310

0.2

Polished (320–600 alumina grit)

596

–

Belted and glass bead blasted

610

1.0

Belted, beaded, shot-peened, and grit blasted

505

0.9

Belted, beaded, and grit blasted

555

0.8

Ultrasonic machined

676

–

Shot peened

531

0.8

Ground

359

0.5

Electrical discharge machined

145

0.2

Ti–5A1-2.5Sn

a The actual reduction factor at long fatigue lives, specifically at Nf = 106 to 107 cycles or greater,

is called the fatigue notch factor and is denoted Kf [110]

empirical estimates are commonly used to obtain Kf values which are used in design of orthopaedic replacements [110]. Hosseini et al. [106] concluded in their studies that the small rate of fatigue notch factor reduction due to increase of tensile strength, suggest a good indication of the nonsensitivity of this material to a notch which means that Ti–6Al–4 V is an ideal alloy for complex shape devices which can be used as body components. In spite of the fact that no standard testing for fatigue evaluation of biomaterials has been established, a variety of testing conditions have been reported in different fatigue studies of orthopaedic materials. [109], However, by the early 1980s, the finite element analysis (FEA) that was first introduced to the field of orthopaedic biomechanics in the early 1970s in order to evaluate stresses in human bones, had been established as an effective and reliable simulation tool for evaluating wear, fatigue, crack propagation, and so forth, and moreover, it was used in many types of pre-operative testing. Judging by these advances, it is well accepted that FEA will continue to contribute efficiently to further progress in the design and development of metal and ceramic orthopaedic implants, along with the understanding of other complex systems of the human body [112]. Very recently, Delikanli and Kayacan

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[113] reported interesting results of their research aimed at the design, analysis, manufacturing and fatigue test processes of hip implants using FEA (in accordance with the requirements of ISO 7206-4:2010 standard) obtaining a satisfactory fatigue performance enabling implants manufactured from Ti–6Al–4 V alloy to be used in THA. They created semispherical pores on the side surfaces of the implants with diameters varying from 0.3 mm to 1 mm, with 0.1 mm increments, for a total of nine specimens (three of each 0.3 mm, 0.6 mm, and solid) using a direct metal laser sintering (DMLS) machine. The obtained specimens were classified as solid specimens (T1, T2, T3), large pore (0.6 mm) specimens (BG1, BG2, BG3) and small pore (0.3 mm) specimens (KG1, KG2, KG3) as well as a reference implant (O) produced by machining. Remember, that in accordance to the fatigue test standard, the implants must withstand loads during 5 million cycles without any damage to be able to successful. Table 8.5 presents an excellent collection of results corresponding to the afore mentioned fatigue tests, which in turn suggests that DMLS can be a suitable alternative for manufacturing implants with sufficient fatigue performance compared with machined ones. Additionally, implants with increased pore diameters caused an approximately 3.82% higher displacement (vertical displacement of the implants which was measured via a displacement sensor connected to an actuator in the test machine at every 50,000 load cycles) values whereas the solid implants showed a lower displacement compared to the lightened ones, corroborating that the solid implants have a higher stiffness than the lightened ones. Within the different fatigue test methods for orthopaedic implants, FEA is increasingly being used in the design, development, and preclinical testing of joint prostheses. However, it is necessary to ensure that these future studies embrace a wider criterion involving all possible parameters that affect the fatigue performance of metallic and/or ceramic orthopaedic prosthesis and incorporate patient-specific variabilities [114]. In other words, Table 8.5 Mean displacement values of specimen groups after 5 × 106 load cycles. From Ref. [113]. Delikanli YE, Kayacan MC. Design, manufacture, and fatigue analysis of lightweight hip implants. J Appl Biomater Funct Mater 2019;17(2):1–8. Under Creative Commons Attribution-Noncommercial 4.0 Licence © The Author(s) 2019

Specimen

Fatigue test displacement (mm)

Mean displacement (mm)

FEA displacement (mm)

O

0.235

0.235

–

0.228

0.22451

0.277

0.24214

0.288

0.25308

T1

0.170

T2

0.273

T3

0.240

KG1

0.280

KG2

0.225

KG3

0.327

BG1

0.191

BG2

0.346

BG3

0.328

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simulations must include holistic and closely corroborated, multi-domain analyses which account for real world variability and so provide useful advantages to complement laboratory experiments. It is important to be aware that there are a limited number of reports on the systematic investigation and comparison of fatigue characteristics, and also a shortage of reports related to fatigue characteristics obtained in vitro or in vivo, as well as a scatter of the fatigue data among the various published studies [90]. Undoubtedly, the fatigue characteristics of metallic biomaterials should be further studied, and reliable fatigue data should be accumulated. In this context, the reader is referred to consult Ref. [101] for a more detailed and interesting study of the fatigue strength of biomedical stainless steel, Co alloys, titanium and its alloys, and bone.

8.1.4

Corrosion

Corrosion has been a persistent challenge in orthopaedics that even Sir John Charnley (pioneer of THA and creator of the “Wrightington centre for hip surgery”) noted as a critical challenge in the design of arthroplasty implants. This can lead to numerous problems such as implant failure from loss of mechanical integrity, third-body wear owing to trapped particles between the articulating surfaces, osteolysis, and localised granulomatous reactions [115]. However, despite more than four decades of experience, the phenomenon of corrosion remains a problem in orthopaedic implants. Nowadays, biomaterials are commonly made of metals and alloys, ceramics, polymers, and composites. Therefore, corrosion is an important factor to take into account in the design and selection of metals and alloys for service in vivo inasmuch as allergenic, toxic/cytotoxic or carcinogenic (for example Ni, Co, Cr, V, Al) species may be released into the body during corrosion processes leading to implant loosening and failure [63, 116–122]. Considering this problem, it is necessary that corrosion tests and/or the solubility of candidate materials are determined before they are approved by regulatory organisations. According to Park and Lakes [33] an optimum metallic material used in THA must have great biocompatibility, excellent resistance against corrosion and wear, an acceptable strength to endure the cyclic loading, as well as a low Young’s modulus in order to minimise stress shielding to the bone. With corrosion being particularly important, in vivo, this depends on the material, the mechanical and geometric properties of the implant, and to a great measure, on the surrounding environment [123] since the behaviour of the implant is different once implanted compared to its behaviour in air [4]. In fact, the longevity of an implant depends mainly on the pH factor that differs from one human body to another, which in turn can temporarily change due to an injury and/or inflammatory process or as a result of corrosion, according to the investigations of Kuncická et al. [11]. Keeping in mind that most metals can generate a passivating layer, their corrosion resistance can be improved by coatings and enhancement of surfaces of the implanted components with other elements such as N2 ,

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O2 [124]. In spite of this, some corrosion resistant alloys (for example stainless steel) can cause allergic or inflammatory reactions, especially in long-term implantations. Within the most predominant types of corrosion in THA, the most relevant ones are galvanic corrosion, pitting, crevice corrosion, and mechanically assisted corrosion [125]. It is important to emphasise that galvanic couples between dissimilar metals are the source of the most common form of corrosion [126] and can be avoided in orthopaedic implants by using alloys that have passive protective oxides, or that are close to each other in the galvanic series [11]. In reality, neither surface coatings nor passivation oxide layers can prevent galvanic corrosion completely considering that the presence of pores and possible surface cracks can lead to direct contact of the corrosive media with the substrate. The corrosion properties of orthopaedic alloys can be evaluated by a wide variety of methods involving immersion-testing, electrochemical tests (anodic polarisation and linear polarisation), and other specialised tests such as impedance spectroscopy and stripping analysis. These methodologies are mainly aimed at investigating the rate of ion release, the electrochemical conditions that cause oxidation and reduction processes, and the electrical nature of the interface [127, 128]. A full discussion of the types of corrosion in hip implants as well as corrosion test methods is beyond the scope of this review, and these themes are available elsewhere [117, 118, 120, 124–129]. It is well known that the metals exhibit high strength and toughness, but they also have several drawbacks such as high modulus of elasticity, E (which might cause stress shielding of the bone) and high susceptibility (usually increased by the action of applied forces and wear) to chemical and electrochemical degradation [117]. As mentioned previously, bioinert metals such as stainless steel (316L), cobalt-chromium (CoCr) alloys and titanium (Ti) alloys are the most commonly used metals in orthopaedic arthroplasties, fixation and bone remodelling, taking advantage of their long-term stability in severe reactive in vivo conditions [40, 130]. However, not all are advantages for these bioinert materials since they may lead to material degradation associated with release of unwanted metallic ions causing local tissue damage and inflammatory reactions, gradual osteolysis which may undermine the fixation, and ultimately, the loading and force transfer [131]. The use of medical grade stainless steel has been a common practice within the orthopaedic community from many years [132, 133] because it possesses advantages in mechanical properties, has an ease of manufacture, and is widely available at low cost [134, 135]. The corrosion performance of stainless steels has been investigated both in vitro and in vivo in numerous papers. These have reported results on various topics such as the effect of antibiotic additions to saline on the corrosion potential of the major surgical alloys based on stainless steel [136]. The amorphous oxides on 316L stainless steel were resistant against localised corrosion both in vitro and in vivo, whereas their counterpart polycrystalline oxides showed severe pitting or crevice corrosion [137]. The in vitro crevice corrosion of cold worked 316LVM (vacuum melted (VM)) steel could be reduced through passivation in 30% HNO3 [138]. Cyclic anodic polarisation tests with highly loaded fracture mechanics samples resulted in a lower Eb suggesting a lower resistance to

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pitting corrosion [139], and there was a tenfold decrease in fretting corrosion when 10% solution of foetal calf serum was added to saline [140]. Zitter [141] suggested a pitting resistance equivalent number PREN of 26 and greater in order to prevent in vivo pitting corrosion of stainless steels, in comparison to the value of 40 usually required for stagnant seawater. Likewise, additions of nitrogen to 316L stainless steel increase the PREN of this alloy to above 26, and also increased the UTS at the expense of some loss of ductility (as observed by elongation to fracture) [117]. Kamachi et al. [121] identified a variety of failure mechanisms in stainless steel orthopaedic devices and suggested alternative routes to improve the corrosion behaviour of these steels. Alloy modification was carried out through additions of titanium and nitrogen as alloying elements, resulting in a super-ferritic stainless steel (Sea-Cure., UNS S44660, ASTM A268 [142] with high strength and high resistance against localised corrosion, and another stainless steel (317L (UNS S31703) which was expected to have a significantly higher resistance to pitting and crevice corrosion than 316L (austenitic stainless steel used in hip replacement containing around 17% Cr, 12% Ni and 2.5% Mo with carbon below 0.03%) at ambient temperatures, when the nitrogen contents of 680–1600 ppm were added [143]. From these investigations, it is apparent that increasing the nitrogen concentration in austenitic stainless steels leads to higher values of Eb that indicate a higher resistance to pitting corrosion. Based on results regarding the resistance of the different stainless steels to pitting corrosion, Kamachi et al. [121] ranked them as follows: super-ferritic > duplex > 316L with 1,600 ppm nitrogen > 317L with 1,410 ppm nitrogen > 317L with 880 ppm nitrogen > 316L with 680 ppm nitrogen > Ti-modified 316L > reference 316L. Meanwhile, with respect to crevice corrosion resistance, the order was slightly changed, as follows: super-ferritic > duplex > 317L with 1,410 ppm nitrogen > 316L with 1,600 ppm nitrogen > 317L with 880 ppm nitrogen > 316L with 680 ppm nitrogen > Ti-modified 316L > reference 316L. Advanced stainless steels for surgical implants include the lownickel (or nickel-free) austenitic stainless steels [10], an example of which is UNS S29108 [144], commercially known as BioDur® 108 from Carpenter Technology Corporation, which was developed to minimise problems associated with Ni toxicity and is produced following the electro-slag re-melting (ESR) process with the purpose to guarantee its microstructural integrity and cleanness. This alloy has better mechanical properties (both static and fatigue) and an excellent localised corrosion resistance compared to type 316L alloy [117]. It is important to note that environment inside the human body is both physically and chemically different from ambient conditions and therefore, a stainless steel replacement that performs well (is inert or passive) in the air could suffer severe corrosion in the body. However, most corrosion resistant stainless steels for orthopaedic applications can cause chronic allergies and toxic reactions in the host body, which are regrettably only diagnosed after a sufficiently long post-implantation period [4, 145–147]. Previous studies in the literature have reported eight separate types of corrosion where only a few have

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8 Mechanical Aspects of Implant Materials

a major impact on stainless steel manufactured by different technologies and under different corrosion tests. The main results regarding to above mention are summarised in Table 8.6 [148]. From Table 8.6 it is very interesting to note the great variety of results concerning the corrosion of metallic implants manufactured with stainless steel which makes it extremely challenging to simulate the exact in vivo environment in a laboratory setup. According to Chen and Thouas [4], the corrosion resistance of stainless steel and other metallic implants must be minimised in the harshest conditions of the body in order to remain at a satisfactorily low level over a long service period (more than 30 years) under normal physiological conditions. With this in mind, owing to their high Cr content and the formation of a Cr2 O3 surface oxide layer, the corrosion resistance of CoCr-based alloys is far greater than for stainless steels [4, 11], including for chlorideinduced crevice corrosion. Moreover, galvanic corrosion is also of less concern compared to stainless steels [117] however, their surface layer is prone to pitting corrosion in Cl-rich surroundings [149]. Based on these scenarios, the Co–Cr alloys are an excellent alternative for orthopaedic implants in THA. CoCr alloys with additions of molybdenum have been one of the load-bearing materials of choice for joint replacement surgery as indicated by its clinical usage over the last forty years as THA and TKA [166, 167], since Mo in solid solution can refine the crystal grains and improve the tensile strength [168]. Nevertheless, when this alloy is implanted into the body, the MoM joint will experience tribo-corrosion, a form of metal degradation [169]. There is an imperative need to study the tribo-corrosion of MoM inasmuch as for the case of Co–Cr–Mo alloys, they produce a significantly higher release of metal ions in comparison with MoP bearings [170]. On the other hand, Co–Cr–Mo alloys are the most commonly used MoM bearing due to their high corrosion and wear resistance [171], properties resulting from the addition of carbon which results in the formation of carbides in the microstructure. Typically, 5–7 wt. % molybdenum is used to improve the mechanical properties of the alloy as it provides solid solution strengthening and good localised corrosion resistance [169, 172, 173]. Likewise, Co–Cr–Mo alloys with high carbon have been shown to have superior corrosion and tribo-corrosion resistance leading to a reduction of low carbon products as bearing materials in the body [173]. It is known that one of the main underlying causes of pain, implant loosening, and/or implant failure results from localised corrosion [167] and there are excellent theories regarding the causes of corrosion as well as the various resources to determine which factors are prone to accelerate corrosion. However, investigators, surgeons, and research physicians view corrosion from a clinical perspective, and they have included elevated metal ion concentrations in synovial fluid and blood, and pseudotumor and necrotic black tissue around implants, as some of the clinical concerns of a corroding implant. Undoubtedly, factors such as manufacturing conditions, tolerances and operator skill during implantation are additional variables to the potential causes of implant failure, where the corrosion of this implant, is a phenomenon that potentially affects the outcome of all patients with implants [174]. In light of this, no metal implant can avoid corrosion [167].

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Table 8.6 Principal results about corrosion studies of stainless steels. Adapted from Ref. [148]. Méndez CM, Covinich MM, Ares AE. Resistance to Corrosion and Passivity of 316L Stainless Steel Directionally Solidified Samples, Chapter 3, in Developments in Corrosion Protection, Edited by M. Aliofkhazraei, IntechOpen, (February 20th, 2014), pp. 41–63. Under the terms of the Creative Commons Attribution Licence. © 2014 Méndez et al.; Licensee InTech Date References Summary of main results 2001 [150]

The presence of Mo assists the rapid re-passivation of the bare metal so that further metal dissolution and, therefore, metastable pit formation is prevented

2002 [151]

A depressed capacitive loop at high frequencies indicating a charge transfer process; a second capacitive loop at intermediate frequencies attributed to adsorption/desorption processes; and a third capacitive response at low frequencies associated with a diffusion process through corrosion products inside pits

2002 [152]

The presence of Widmanstatten, intergranular austenite and partially transformed austenite exhibited beneficial effect on SCC resistance by deviating the crack propagation path

2004 [153]

Crevice corrosion behaviours of stainless steels containing 25% Cr, 3% Mo and various amounts of Ni were investigated by immersion tests in natural seawater and in a 10% ferric chloride solution. In the seawater immersion, 4% Ni ferritic steel showed higher resistance than 0% Ni ferritic steel and 30% Ni austenitic steel. In contrast, the CCT in 10% ferric chloride solution was steadily decreased with increasing the Ni content

2005 [154]

Improvement of the re-passivation of the passive layer observed in this group has to be related to the increase of Mo concentration in the passive layer

2006 [155]

Increasing Nb content results in increasing the localised corrosion resistance of austenitic stainless steels in NaCl. Cold deformation (CD) has a critical role on the corrosion resistance of stainless steel

2008 [156]

The MnS inclusions do not undergo passivation; thus, a large density of locations susceptible to pitting attack in a chloride-containing environment is available

2008 [157, 158]

Molybdenum seems to play, therefore, a dual role in the improvement of the general corrosion resistance of the stainless steels by modification of passive film composition and modification of active dissolution by formation of insoluble oxides

2008 [159]

The carbon steels 304 SS and 316L SS have been markedly affected by water line corrosion

2009 [160]

The structure of the passive film is potential dependent and affects the efficiency of the cathodic protection current

2010 [161]

The re-passivation rate of Fe–18Cr–xMn (x = 0, 6, 12) was significantly reduced with Mn content. The susceptibility to metastable pitting corrosion was significantly increased with Mn content (continued)

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8 Mechanical Aspects of Implant Materials

Table 8.6 (continued) Date References Summary of main results 2010 [162]

Three regions, namely, the passive, active (consisting of severely attacked, commonly attacked, and lightly attacked regions) and variable regions, can be observed on most crevice corrosion sites

2011 [163]

The microhardness was not increased after melting treatment, but corrosion resistance improved especially when a high power is used. There was a shift in Icorr toward lower and nobler values after this treatment

2012 [164]

The films formed on AISI 316 under DC. potentiodynamic polarisation are thicker, since the metallic contribution was hardly, or could not be detected. Moreover, the outer layers of the cycled films are extremely rich in Fe3+ oxides and depleted in Fe2+ , chromium oxides and nickel oxides

2012 [165]

The evaluation of marine crevice corrosion mainly through the investigation of re-passivation potentials is the best approach to characterise the resistance of the alloys to localised corrosion

In order to obtain customisable and patient-specific implants, advances in technology over the last ten years has allowed the manufacture of Co–Cr–Mo implants via additive manufacturing (AM) methods with relative ease, reaching a fine and more dispersed carbide structure [175, 176] with an improved tribo-chemical performance. Indeed, with the rising adoption of AM technologies and the imminent risks associated with MoP [177, 178], there has been an increase in research focussed on the tribo-corrosive behaviour of biomedical alloys, particularly Co–Cr–Mo alloys. Nonetheless, the degradation of metals caused by mechanical motion and corrosion remains a potential concern [129, 179–183]. As previously mentioned, Co–Cr–Mo alloys form a protective oxide (Cr2 O3 ) layer whose nature is determined by its composition, structure, thickness, and the amount of defects present. Additionally, its protectiveness can be altered by environmental factors such as pH, chloride ions, exposure duration, and so forth. On the other hand, the corrosion behaviour of such alloys is influenced by several factors including material characteristics (chemical composition, microstructure, surface features), external environment (pH, temperature, oxygen content), as well as manufacturing conditions, for example, cold working and thermal treatments [167, 184]. The human body is a highly corrosive environment for metals and the body´s fluids (human plasma, synovial fluid of joints) are composed of 0.9% sodium chloride (NaCl) with additional inorganic salts and various protein molecules [184]. These fluids are highly corrosive in nature as a result of the high concentration of chlorine ions in solution, which can accelerate pitting corrosion. Additionally, observed crevices between components joint prosthesis, and inflammation after prosthesis implantation, can result in significantly lower pH values (pH < 5) that can accelerate crevice corrosion [167]. Therefore, it is important to be mindful of the different variables that can alter the corrosion behaviour of Co–Cr–Mo alloys.

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Despite their applications in orthopaedics, the electrochemical properties of Co–Cr–Mo alloys have not been extensively reported until now [185, 186]. According to Bettini et al. [187], most studies focus on the influence of the microstructure on the wear and mechanical properties of Co–Cr–Mo alloys, but little is known about its influence on the corrosion behaviour of Co–Cr–Mo alloys. Additionally, it is important to state that there is a certain difficulty in comparing the results from different research works since the different experimental parameters such as temperature, electrolyte composition, time of immersion and potential in which the electrochemical measurements were done, amongst others, can vary notably. A recent investigation [185] on corrosion behaviour of an orthopaedic implant of Co–Cr–Mo alloy exposed to the physiological serum and Hank’s solution under different temperatures and electrolyte compositions, as well as immersion time, were simulated by means of alternative methods such as open circuit potential (OCP), potentiodynamic polarisation curves (PPC), and electrochemical impedance spectroscopy (EIS). The authors concluded that Co–Cr–Mo biomaterial is greatly influenced by electrolyte composition and temperature, and moreover, the passive films were more protective when formed in the presence of a physiological serum. On the other hand, EIS results suggest that the layer formed in the presence of physiological serum is nobler than that with Hank’s solution. Keeping in mind that in modular THA, the micro-motion leads to fretting corrosion in the head/neck and neck/stem interfaces, this being the major cause of early revision in hip implants and more specifically, those with heads larger than 32 mm, Kofi [188] found that the type of fluid used to simulate the fretting corrosion of Co–Cr–Mo biomaterials is crucial for the reliability of laboratory tests and therefore, the need to replicate the human body environment as closely as possible during in-vitro testing and validation, is of paramount importance. On the other hand, the problem with MoM couples (Co–Cr–Mo implants) in THA is the tribo-corrosive behaviour of this alloy whose long-term performance depends greatly on the overall tribo-corrosion behaviour of the material and not just on the individual contribution from wear or corrosion, according to Krasicka-Cydzik et al. [189] and Toh et al. [175]. Based on the above, it is also important to consider the effects of tribo-corrosion on Co–Cr–Mo alloys under the different conditions that undoubtedly play an influential role in the tribo-corrosive behaviour of the Co–Cr–Mo alloys [175, 190]. Metal carbides, effects of alloying elements, effects of tribo-corrosion conditions, influence of wear debris, and influences of manufacturing methods are the main parameters that supposedly have a great influence on the tribo-corrosive behaviour of such alloys, and these can be consulted in more detail in Ref. [175]. Figure 8.2, illustrates a summary of the above-mentioned parameters which have an influence on the tribo-corrosive behaviour of Co–Cr–Mo alloys, where the green arrows represent the interactions of these different parameters. From the practical point view [175], the OCP and EIS techniques are convenient and well-accepted methods to quantitatively evaluate the tribo-corrosive properties of the concerned alloy. However, methods such as potentiostatic and cathodic protection must not be ignored as they provide important means of evaluating the tribo-corrosion process kinetics occurring between the alloy and the environment

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Fig. 8.2 Schematic illustration of the five types of influence affecting the tribo-corrosive behaviour of Co–Cr–Mo alloys. From Ref. [175]. Toh WQ, Tan X, Bhowmik A, Liu E, Tor SB. Tribo-chemical characterisation and tribo-corrosive behaviour of Co–Cr–Mo alloys: A review. Materials 2018;11:30. Under the terms and conditions of the Creative Commons Attribution (CC BY) licence. Licensee MDPI, Basel, Switzerland. © 2017 by the authors

during wear since the long-term performance of Co–Cr–Mo alloys largely depends on the applied environmental conditions. A more clinically relevant simulated body environment [188] must be widely investigated and reported in order to understand the fundamental corrosion behaviour of Co–Cr–Mo alloys in THA. In comparison to other biomaterials such as stainless steels and cobalt-based alloys [191–195], nowadays commercially pure Ti (CPTi) and Ti alloys are being extensively used in orthopaedic applications as implant materials taking advantage of their low density, high specific strength, and excellent corrosion resistance as well as high biocompatibility. It is worth noting that for orthopaedic applications, the combined effect of corrosion and friction is of paramount importance emphasising that in the case of titanium and its alloys, this corrosion mode is dependent, in a great measure, on their crystal structure [193, 196]. Like all metallic implants, as a consequence of degradation, these release metal products into the periprosthetic tissue [197] provoking inflammatory events, commonly termed adverse reactions to metal debris (ARMD), and this is associated with MoM hip resurfacing (HR) implants with Co–Cr–Mo bearing surfaces [198]. However, a notable number of failures secondary to ARMD in a Co–Cr–Mo/Ti interface, between the neck and stem junction where the MoM bearing surfaces are absent, have been reported [199– 201]. This modularity at the head-neck junction of stems in THA became very popular in the 1980s [202], simplifying subsequent revisions since there is the option of retaining

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the stem and performing a head exchange [200]. Nevertheless, in spite of these potential benefits, it is observed that an increase in the use of modular interfaces inevitably leads to an increase in fretting and crevice corrosion at the taper junction [203–209], and the confirmation of corrosion at the femoral head neck junction is made at revision surgery. A device as such is described by Hussenbocus et al. [202]. The Oxford English Dictionary defines the word “taper” as “reduce in thickness towards one end. A more descriptive definition is the uniform change in diameter of a cylindrical object measured along its axis. The concept of a Morse taper is that of a cone within a cone-the trunnion (male portion) and the bore (female portion)-and the stresses created by the compression of the wall of the bore by the trunnion causes an interference fit (Fig. 8.3) and cold-welding between the two components, which increases during physiologic loading. The radial compressive stress causing friction between the taper surfaces also provides resistance to separation at the head-neck junction. The original Morse taper cone angle described was 2◦ 50´ but within orthopaedics, the term Morse taper is loosely used to encompass tapers of all angles that result in cold-welding of one element upon another”. The head-neck junctions of femoral stems are usually composed of Ti–6Al–4 V alloy or CoCr alloy, where both form a protective surface oxide layer through a process of self-passivation, favouring corrosion resistance. The presence in the body of mixed metal systems, such as Ti–6Al–4 V stems with CoCr heads, can induce galvanic corrosion which can cause more loss of material from the bore (B in Fig. 8.3) of the head than from the trunnion (T in Fig. 8.3) [200, 204–206]. On the other hand, titanium alloy is not a commonly used material as a femoral head due to its lower modulus of elasticity and inferior wear properties compared to Co–Cr or ceramics [116, 210]. An example of this is shown in Fig. 8.4 [202]. The contribution of fretting and corrosion damage at

Fig. 8.3 Diagrammatic representation of a femoral head-neck taper junction consisting of a female bore (B) and male trunnion (T). The trunnion length (TL), proximal cone diameter (PCD), distal cone diameter (DCD), and cone angle (CA) are displayed. From Ref. [202]. Hussenbocus S, Kosuge D, Solomon LB, Howie DW, Oskouei RH. Head-neck taper corrosion in hip arthroplasty. BioMed Res Int 2015: Article ID 758123, 9 pages. Under the Creative Commons Attribution Licence. Copyright © 2015 S. Hussenbocus et al.

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Fig. 8.4 Fretting and corrosion damage on the a bore of a CoCr femoral head and b Ti-alloy trunnion from a retrieved failed metal-on polyethylene (MOP) THA. From Ref. [202]. Hussenbocus S, Kosuge D, Solomon LB, Howie DW, Oskouei RH. Head-neck taper corrosion in hip arthroplasty. BioMed Res Int 2015: Article ID 758,123, 9 pages. Under the Creative Commons Attribution Licence. Copyright © 2015 S. Hussenbocus et al.

the taper junction [116] to implant failure is just being recognised. Thus, the severity of early corrosion-related failure of the modular femoral stem neck junction is a potential reason for concern [211], and in this context, the surgeon must be aware of the potential drawbacks of modularity in THA [202]. Clinical flexibility is one important reason for the use of modular designs of the hip implant, where the stem and the ball are commonly made of two different materials. According to the literature, a common practical solution is the ball being made of highly polished Co–28Cr–6Mo alloy or a ceramic composite such as alumina with additions of zirconia, whereas the stem is made of Co-based alloy [212–214]. However, this type of modular hip prosthesis is susceptible to both galvanic and fretting corrosion, as well as related problems, mainly at the taper interfaces as a consequence of the micro motion, crevices at the taper mismatch, and galvanic coupling of dissimilar materials, or a combination of all these three components, as reported by Choroszy´nski et al. [212]. However, it has been shown that ceramic heads greatly mitigate fretting corrosion, according to data by Kurtz et al. [215]. It is curious that despite the many advantages of titanium alloys they have tendency to stick which leads to poor friction and abrasive properties. Hence, the sliding contact of these alloys should be avoided unless the surface layers are modified by thermochemical treatments such as nitriding and/or carbonitriding, for example [216]. Moreover, it is important to point out that fretting damage in titanium alloys significantly reduces the fatigue life with fretting being the major concern for hip implants in joints between femoral stems and heads, since fretting corrosion is the main mechanism of gradual release of metal ions to different tissues in the body causing pain and accumulation of wear particles resulting in the prosthesis loss, and in some cases, adverse biological reactions [212, 217]. An interesting study by Okazaki and Gotoh [218] related with the release of metal ions from the Ti–15Zr–4Nb–4Ta alloy in pseudo body fluids in comparison with those from

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Ti–6Al–4 V and vanadium-free Ti–6Al–7Nb alloys that are widely used as implantable titanium alloys throughout the world, revealed promising corrosion resistance results. According to Okazaki and Gotoh [218], the quantities of the titanium ions and each element released from the three titanium alloys studied, changed depending on the element added to the titanium alloy as well as the composition of the immersion test solution. In fact, the total quantity of zirconium, niobium and tantalum ions released from the proposed Ti–15Zr–4Nb–4Ta alloy was much smaller in comparison to that of elements released from the Ti–6Al–4 V and Ti–6Al–7Nb alloys. With these results, the authors [218] manufactured, amongst other implants, an artificial hip joint with the Ti–15Zr–4Nb– 4Ta alloy using conventional manufacturing processes and obtained an excellent corrosion resistance with great expectations for implant applications in the future. It is noteworthy that part of the neck on the stem surface of the artificial hip joint was experimentally fabricated by the same surface treatment used for advanced cementless artificial hip joints made of the Ti–6Al–2Nb–1Ta alloy, following procedures reported by Ido et al. [219] and Yamamuro et al. [220]. Regardless of the type of titanium alloy, their corrosion resistance is strongly dependent on the resistance of the oxide covering the surface. This can vary depending on the surrounding environment [221]. However, it is important to keep in mind that an ideal coating must be a system in which anticorrosion, anti-infection, and osseointegration must be obtained simultaneously [222], and for orthopaedic applications, some titanium alloys are very close to this ideal alloy [223]. It is well known that the main goal in orthopaedic surgery is that the used prostheses have to last an entire lifetime immersed in human body fluids without the possibility of inspection and maintenance. For this reason, their corrosion resistance is of paramount importance. In fact, the scientific literature search leads to a great amount of reports concerning the corrosion resistance of titanium alloys in human body applications where the most common solution used to study the electrochemical behaviour of the metal is simulated body fluid at pH 7.4, according to Yingjie et al. [224]. This is reflected in the use of more than 1000 tons of titanium devices, of every description and function that are implanted in patients worldwide every year [225]. As mentioned, the pH is an important factor which varies in different solutions and therefore, every solution may impose a different electrochemical result on a specific alloy, so studying different simulated body fluid solutions would provide a major understanding about components and the biomedical corrosion procedure in these alloys [226]. Liang and Mou [227] studied the corrosion characteristics of Ti–6Al–4 V alloy in different simulated body fluids including Ringer’s solution, phosphate-buffered saline (PBS) solution, and Hank’s solution, under different pH values. They found out that the order of corrosion rate (from highest to lowest) was Ringer’s solution, PBS solution and Hank’s solution, respectively. These results have been corroborated by many authors, for example [228, 229], indicating that the Ringer’s solution seems to be more aggressive due to more concentrations of Cl− ions. An outstanding investigation about the effect of different ions on the corrosion rate of the tested alloys for each simulated body fluid solution is reported in [230]. Admittedly, chemical composition

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is a parameter which influences the corrosion resistance of titanium alloys including new beta type titanium alloys. For instance, Gebert et al. [231] reported that small additions of indium (In) to a β-type Ti–Nb alloy and tested in Ringer’s solution does not have a negative effect on the corrosion behaviour, but further additions of indium can produce an enhanced corrosion resistance of the alloy. The corrosion components lead to aggressive media for biomedical implants implying that the control and improvement of the corrosion resistance of orthopaedic implants is linked to the manufacturing process of such implants [226, 232]. Recently, β-type Ti alloys containing Nb, Zr, Ta, Mo, Sn, etc. have been produced with an elastic modulus significantly reduced by adjusting the concentration of the β stabilising elements [233– 235]. These types of Ti alloys demonstrate an extraordinary corrosion resistance in human body fluid as a consequence of the formation of a hard and tightly adherent protective oxide film [236]. Therefore, these Ti-based alloys with nontoxic and non-allergic elements are being widely used to design new β-type Ti alloys [20]. It is important to keep in mind that there are various groups of new generation titanium alloys and therefore, they may offer different levels of corrosion resistance in the body fluid depending on factors such as alloying elements and their amount, as well as fabrication methods. The importance of this subject and the increasing research interest are the driving force for the development of superior β-Ti alloys for orthopaedic applications [234]. It has already been reported that [237–239] even for two distinct alloys with the same chemical composition but different fabrication methods, the corrosion and electrochemical behaviour may be different. Likewise, alternative heat treatments as well as mechanical processes of Ti alloys involving ageing and cold working could have great influence on the microstructure of the alloy and consequently, the corrosion behaviour. With regards to the novel multifunctional β-type titanium alloys applied to biomedical fields [240–243], the proposed Ti–24Nb–4Zr–8Sn (by weight percent) alloy exhibited an excellent compromise between a low elastic modulus and high strength, as compared to previously reported alloys [244, 245]. It was observed that the passive region of the Ti–24Nb–4Zr–8Sn alloy was wider than those of the commonly used Ti–6Al–4 V and pure titanium alloys. On the other hand, the corrosion current density for this β-type titanium alloy was as low as that of pure titanium [240]. Lin et al. [246] carried out a study on the corrosion behaviour and the effects of solution treatment (ST) and ageing on both the microstructure and mechanical properties of a β type alloy Ti–40Ta–22Hf–11.7Zr (denominated as TTHZ) manufactured by the coldcrucible levitation technique. The solution-treated TTHZ alloy exhibited a single β phase with the highest corrosion resistance, in comparison to the as-cast and the solution-treated followed by aging samples. The OCP analysis indicates that the corrosion resistance of the as-cast TTHZ alloy was superior to those of both CP-Ti and Ti–6Al–4 V. Therefore, it can be concluded that the order of the corrosion resistance of the three alloys was as-cast TTHZ alloy > CP-Ti > Ti–6Al–4 V, suggesting that the corrosion mechanism of β titanium alloys is attributed to the formation of a compact and stable passive oxide coating

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on the material surface. Likewise, Lin et al. [246] reported that the corrosion resistance of the TTHZ alloy seems to be influenced by the alloying elements Ta, Hf, and Zr inasmuch as investigation about corrosion resistance of Ti–Ta alloys revealed that an increase in Ta content leads to an increased corrosion resistance of the alloy [247]. Alternatively, Zhou and Niinomi [248] demonstrated that thanks to the passive layers of HfO2 and TiO2 formed on the Ti–Hf alloy surface, the corrosion resistance is superior to that of pure titanium. There are many options or possibilities to develop β type titanium alloys that can be an excellent alternative, combining a low elastic modulus and good corrosion resistance, such as Ti–30Ta–10Nb–10Zr and Ti–40Ta–10Nb–10Zr alloys developed by Kim et al. [249], who showed that these alloys can be promising metallic biomaterials for future use in the orthopaedic field. The presence of β-phase in these alloys is fundamental to producing an excellent workability along with an isotropic mechanical behaviour owing to the body centred cubic (bcc) structure [250, 251]. Nowadays, there are many studies on titanium and its alloys for applications in both the aeronautical industry and in orthopaedics. However, and despite the most commonly used alloys in the latter field, CP Ti or Ti–6Al–4 V, the β type alloys are the ones that undoubtedly provide excellent corrosion resistance [14]. However, between the variety of β titanium alloys, the most important commercial and experimental ones are Ti–29Nb– 13Ta–4.6Zr (wt. %) alloy and the Ti–13Nb–13Zr (wt. %) alloy. Nevertheless, the Achilles heel of these alloys is their high costs associated with extraction, processing, and alloying and therefore, their applications are very specific. Furthermore, they also possess higher densities than other classes of Ti alloys due to the high concentrations of beta stabilising elements, which is a disadvantage when low mass is a requirement. New β alloys remain a challenge in order to achieve commercial success in biomedical implant applications and it seems that the β typeTi–35Nb–7Zr–5Ta alloy with a porous structure, recently developed for such applications by means of powder metallurgy, is the most suitable choice for the successful osseointegration [252]. The studies on the corrosion of Mg alloys for biomedical applications cannot be ignored in spite of the fact that stainless steel, Co–Cr alloys and Ti and Ti alloys are the most commonly used materials in orthopaedics. Considering that these metals corrode gradually with an appropriate host response releasing corrosion products, it is of paramount importance to use essential metals metabolised with a nontoxic effect. In this context, magnesium serves this aim best since it plays an essential role in body metabolism and should be completely excreted within a few days after degradation, according to Pogorielov et al. [253]. On the other hand, there is the technological possibility to control their high corrosion rates, and hence these alloys have a significant opportunity as temporary bio-implants [254]. However, and in spite of several advantages of the Mg alloys such as biocompatibility and osteogenesis [5], biodegradability and avoidance of second surgery [255], favourable mechanical properties [256], machinability and dimensional stability [257], and a high damping capacity [5], the exceptionally high

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corrosion rate in the physiological environment remains the major drawback of magnesium implants [258–261]. This, in turn, promotes the fast degradation of the mechanical properties which leads to the premature failure of implants before the tissue has sufficiently healed, and the production of hydrogen gas in the corrosion process at a rate [254] that is too fast to be dealt with by the host tissue [5, 262]. It is speculated that in spite of some early successes with Mg-based implants [263], this metal was abandoned when stainless steels became available due to the production of gas during the in vivo corrosion process. While a substantial number of reports generate intrigue regarding the use of magnesium and its alloys in orthopaedic implants for load bearing applications, for example THA [5], a lot of research is still necessary to appropriately evaluate magnesium’s potential. This can be through the use of high purity magnesium or alternatively, experimenting with diverse alloying elements but taking care that due to the solubility of alloying elements in crystalline Mg being limited, the corrosion rates can only be altered within a limited range [4]. Moreover, surface treatments are another possibility to improve the corrosion resistance, processes that of course must lead to a non-toxic, biologically compatible material. Up to now, in vitro, and in vivo studies have been used for evaluation of the degradation rate and host response, but regrettably, there is no correlation between these methods, and they should be used together for better alloy assessment. According to Pogorielov et al. [253], “the best method for in vitro degradation is immersion in medium that simulates body environment, such as simulated body fluid (SBF), minimum essential medium (MEM), or Earle’s balanced salt solution (EBSS)”. Nonetheless, the results of in vivo research depend in great measure on the animal species, the implant anatomical location, as well as parameters such as pH, blood flow speed, and chloride ion concentration that can influence or significantly change the corrosion rate and host response. The truth is that there are a few different approaches to improve the corrosion resistance of magnesium implants. The most important of them is the surface modification by conversion coatings including polymeric coatings, calcium–phosphate coatings, and graphene coatings. However, the scientific literature on the corrosion behaviour and the performance of the coatings synthesised on the implants is very scarce. Therefore, and being realistic, the use of magnesium alloys as temporary implants will significantly depend on the in-vivo performance evaluation of these various biodegradable coating systems [254].

8.2

Ceramics

Ceramics have been widely used in biomedical applications for load bearing implants and in the dental industry [264] due to their excellent properties such as hardness, tensile strength, wear and scratch resistance, and mainly, their excellent long-term biocompatibility as a result of their high chemical stability and resistance to corrosion [265]. More specifically, these ceramic bearings are used for the manufacture of femoral heads or acetabular liners for THA where the above properties give orthopaedic implants with a

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low generation of wear debris. However, from a historical perspective, their limitation for orthopaedic applications was also as a consequence of their brittleness, which led to catastrophic failure in vivo [264, 266]. In comparison to the most commonly used biomaterials in orthopaedics, i.e. steel or Co–Cr alloys, the ceramics possess lower density and significantly higher resistance to various corrosive factors. The main mechanical properties are summarised below.

8.2.1

Hardness

The two main ceramics which are widely used in THA today are alumina and its alloys inasmuch as compared to other biomaterials, these offer the advantage of an outstanding hardness making them the ideal materials for the manufacture of components for important applications in the field of orthopaedics [267] where they have proven to be very effective, especially in younger and more active patients [268]. In fact, ceramic bearings made of ZTA have demonstrated the lowest in vivo wear rates to date of any bearing combination considering that the same principles of friction and lubrication reported for MoM bearings apply to CoC bearings [46]. On the other hand, this ceramic oxide has a greater hardness than metal and therefore, can be polished to a surface roughness of a few nanometres. Additional to this, an inconvenience due to the hardness of ceramics is that the wear characteristics are very sensitive to design, manufacturing, and implantation variables [269]. In spite of this, ZTA on ZTA bearings are considered the standard CoC articulation, condition, supported by clinical use for almost two decades. It can be seen that the small amount of generated ceramic particulates are much less biologically reactive than those produced in MoP articulation, and most important of all is that the incidence of osteolysis associated with use of CoC appears to be minimal or non-existent according to a report by Bierbaum et al. [270]. In addition, the hardness of alumina increases its resistance to scratching, and it is much less likely to scratch than titanium or Co–Cr alloys used also in THA. In fact, from the clinical point view, it is of paramount importance since the alumina can undoubtedly resist third-body wear and is not scratched by retained cement particles or bone [46]. Interestingly, scratch resistance also increases with the phase transformation toughening of zirconia. Thus BIOLOX®delta has a higher scratch resistance than BIOLOX®forte, despite BIOLOX®forte having a higher hardness [271]. For this reason, the increased hardness of ceramic materials is considered advantageous. The pattern of damage by a hard third body in metals and ceramics differs notably. Thus, the high hardness of alumina justifies that in some European countries such as Austria, France, Germany, Italy, and Switzerland, more than 70% of the THA are made using ceramic femoral heads, whereas in Asian countries like e.g. Korea, 72% of THAs have an alumina head. Furthermore, in countries where the ceramics encountered some criticism, the use of this type

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of bearing material is today spreading fast [267]. It is important to note that the materials and testing standards on ceramics in order to obtain hardness measurements are not particularly extensive and therefore, the available test results tend to provide only a limited picture of the total material characteristics. As was mentioned in Sect. 6.3, although Y-TZP is one of the strongest ceramics for medical use [272], it has the great inconvenience that femoral heads zirconia vs alumina and zirconia vs zirconia lead to catastrophic failures [273]. In the same context, the LTD or ageing [274] is another drawback of zirconia in orthopaedics and therefore, it is more frequently used mixed with alumina, taking advantage of the hardness of this and the fracture toughness of the zirconia. From this point of the view, reports concerning to the hardness values of Y-TZP as implants in THA are practically non-existent, focussing more on ZTA, or to a lesser extent, to ATZ composites. However, the Vickers hardness (HV) for alumina has evolved over time due to improvements in the manufacturing processes and raw material quality. Generally, the hardness values reported have been 17.7, 18.6, and 19.6 GPa for the first (1970s), second (1980s) and third (2000s) generation of alumina, respectively [275]. These values are very approximate to those obtained and reported by BIOLOX® ceramics, the trade name for medical grade ceramics from CeramTec GmbH (Plochingen, Germany) [276]. They reported HV values of 20 GPa, 20 GPa and 19 GPa for BIOLOX® (since 1974), BIOLOX® forte (since 1995) and BIOLOX®delta (since 2004), respectively. The latter refers to ZTA composites whose composition includes SrO, Y2 O3 , and Cr2 O3 [277], providing increased fracture strength, excellent wear properties and excellent biocompatibility, which is widely used today by orthopaedic surgeons. Some hardness values obtained with alumina manufactured by different processes and modifying the sintering process are reported in [278–283]. It is noteworthy that the strontia was added to the ZTA composite with the aim of forming platelets inside the matrix which support the toughness of the material, meanwhile yttria was included to control the transformation mechanism of the zirconia phase. On the other hand, according to Burger and Richter [277] the inclusion of chromia would improve the hardness in a ZTA material. Regrettably, a scientific study on this subject reveals a measurable increase in hardness only at a chromia content much higher than that present in BIOLOX® delta [284] but within a limited range, chromia does not influence the hardness of a ZTA composite [285]. Kuntz et al. [286] have published a comprehensive study about this topic and found that the minor content of chromia in BIOLOX®delta did not increase the hardness. Burger and Richter [277] have likely measured a hardness increase due to other factors such as the grain size. ATZ is another alumina–zirconia composite bearing material with a general composition of 80% Y-TZP and 20% Al2 O3 , developed and manufactured as Ceramys® by Mathys Orthopaedics [287] in Germany in 2010. This ATZ ceramic revealed a high flexural strength but with a moderate hardness of approximately 14.7 GPa, which is lesser than the ZTA BIOLOX®delta composite, and very similar to the hardness values ranging from 12.7 to 13.45 GPa for pure zirconia and zirconia with additions of 30 wt.% Al2 O3 according to investigations by Maji and

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Choubey [288]. In addition to the LTD of zirconia, there remain major concerns with ATZ composites related to the variation in manufacturing steps carried out by different manufacturers, which needs to be carefully assessed every time before implantation in the human body [289]. Taking into account that for the manufacture of ATZ composites, as well as most other ceramics in the zirconia-alumina system for orthopaedic applications, the high-performance properties are only achieved by expensive and highly energy-consuming processing, for example HIP technology [290]. As an alternative to the mentioned oxide ceramics, there is the possibility of using the non-oxide silicon nitride ceramic material in the orthopaedic field since this ceramic has been used as bearings and turbine blades for more than 50 years [289]. CeramTec GmbH tested the Si3N4 (standard and different SiAlONs) for THA and they found a high wear and high degradation of the material under hydrothermal conditions in comparison with oxides. In spite of this, Si3 N4 has been considered as a bearing material in THA and TKA due to its Young’s modulus of 300 GPa, fracture toughness ranging from 10 to 12 MPa m1/2 , flexural strength of 1 GPa and moderate Vickers hardness of 12–13 GPa, very similar to that ATZ composites. In the medical field, it has been used with relative success in cervical spacer and spinal fusion devices [291] since 2008, with approximately 25,000 implanted spinal cages [292, 293]. Figure 8.5 illustrates the three starting Si3 N4 ceramics tested by CeramTec GmbH. Likewise, Figs. 8.6, 8.7, 8.8 and 8.9 present results of fracture strength, scratch resistance, ring-on-disc, and impact of hydrothermal ageing on fracture strength, respectively. The results for the mechanical tests showed similar values for oxides and nitrides, apart from the bending strength, which was significantly higher for AMC. For all tribological tests with Si3 N4 the temperature of the medium reached up to 60 °C indicating the high friction of these nitrides. The temperature for AMC against itself remained at 40–45 °C throughout the test. The wear for all Si3 N4 variations was independent of the type of nitride used and showed magnitudes higher than for the AMC combination. Therefore, it can be concluded that the nitride ceramics investigated showed neither a mechanical nor a biological benefit over AMC. The CoC combinations of these ceramics demonstrated high friction and a high wear rate. Considering that there remains some discrepancy between the hip simulator and in vivo studies for some femoral heads, clinical studies might be necessary before confirming the use of silicon nitride as a bearing material for hip replacements. Regardless of the material used in the orthopaedic implant, the role of the hardness of the material is very important inasmuch as it implies low wear rates leading to an extended duration in vivo.

8.2.2

Flexural Strength

Remembering that ceramic materials lack plastic behaviour, they also have numerous useful properties such as high hardness, stiffness (modulus of elasticity), and wear resistance, as well as a corrosion resistance associated with chemical inertness. However, their

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Fig. 8.5 Microstructure of 3 nitrides that have been designed, in order to optimise the strength, the hardness and the toughness, respectively as seen in Fig. 8.7. Courtesy of CeramTec GmbH

Fig. 8.6 Bending strength, fracture toughness and Vickers hardness of different Si3 N4 ceramics. (ZPTA) is an alumina matrix containing homogeneously distributed metastable zirconia particles and “in situ” formed hexagonal ternary aluminate platelets. Courtesy of CeramTec GmbH

Fig. 8.7 Scratch resistance of different Si3 N4 ceramics in comparison with metals and oxides. Courtesy of CeramTec GmbH

low density in comparison with steel makes them suitable for technical and biomedical applications where weight reduction is required [294, 295]. Likewise, the more important limitations of ceramics such as alumina, zirconia, ZTA, and alternatively, silicon nitride for a general use in orthopaedics, lies in their relatively low toughness and high stiffness. Moreover, if ceramics are not manufactured properly, it may come down to reliability

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Fig. 8.8 Ring-on-disc results of various ceramic combinations carried out using a ring-on-disc set up according to ISO6474 for a test period of 100 h each in 25% serum. ZPTA is an alumina matrix containing homogeneously distributed metastable zirconia particles and “in situ” formed hexagonal ternary aluminate platelets. Courtesy of CeramTec GmbH

Fig. 8.9 Relationship between hydrothermal ageing and fracture strength for ZPTA and Si3N4 after extreme ageing at 134 °C, 2 bar and a duration of 150 h. ZPTA is an alumina matrix containing homogeneously distributed metastable zirconia particles and “in situ” formed hexagonal ternary aluminate platelets. Courtesy of CeramTec GmbH

issues. This leads to expensive test procedures for the characterisation of the material [294]. Nowadays, flexural testing is the most common method used in order to measure the uniaxial tensile strength of ceramics and glasses under standard test methods using methods that involve rectangular specimens and cylindrical rod specimens, which are the preferred option in many cases, emphasising that the direct tension tests are practically unfeasible due to the large test piece size, the high costs of specimen manufacture and testing, and to a large extent, the difficulties of gripping without misalignments [296]. In this context, flexural strength is a key value to evaluate the stability of a material that will be used as replacement in orthopaedics. The flexural strength test can generate data for many purposes including material characterisation, materials development, and design. The reader is referred to Ref. [296] to analyse in more detail an excellent summary concerning as to how rods have been tested in the past, identifying key experimental errors and remedies, as the foundation for a new standard test method for ceramics and glasses which in turn are being used in alumina and zirconia for orthopaedic applications. It is well known that the flexural strength of a ceramic material depends on its fracture toughness as well as the size and severity of flaws which are commonly introduced in the

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sample during grinding or machining to reach the final size, and may have a pronounced effect on the flexural strength [296–299]. As pointed out by Quinn et al. [296] “with good machining practice, it is possible to obtain fractures from the material’s natural flaws”. An optimum guideline has been effective in minimising or eliminating grinding cracks in several grades of silicon nitride, alumina, silicon carbides and with good expectations for other ceramics. Indeed, the devices presently manufactured from α-Al2 O3 for applications in orthopaedics are characterised by a very fine-grained microstructure following the (HIP) technology and subsequent sintering at temperatures ranging from 1600 to 1800 °C [300]. In this point, it is very important to stress that alumina is very prone to experimental grain growth during sintering which affects the mechanical properties. Commonly, to limit grain growth during sintering, additions of small amounts of MgO (less than 0.25 wt. %) can lead to obtaining a density near the theoretical with an increase in strength, fatigue resistance and fracture toughness. Therefore, these parameters are a function of grain size [301]. According to Ratner [302], with average grain size of less than 4 μm in high purity alumina, it is possible reach a proper flexural strength and excellent compressive strength as recommended for the bearing balls of hip replacements. On the other hand, it is reported that an increase of grain size to levels higher than 7 μm decreases the mechanical properties by about 20% [302]. It is important to guarantee a low level of sintering aids in order to avoid precipitations at the grain boundaries which lead to a degradation of mechanical properties. For example, manufacture of modern Al2 O3 femoral heads doped with MgO to control grain growth under HIP technology at a temperature of approximately 1250 °C, gives an almost fully dense material with a grain size smaller than ~2 μm, and a flexural strength > 550 MPa [265], which is close to the flexural strength of the third generation of alumina commonly used today in THA (580 MPa) [303]. Roy et al. [304] investigated the effect of the sintering temperature of alumina femoral heads manufactured by cold isostatic pressing (CIP) on 3-point flexural strength and found average flexural strength values of 400, 370, and 350 MPa for sintering temperatures of 1550, 1600, and 1650 °C, respectively, indicating that the flexural strength is decreased with the increase in sintering temperature as well as grain size. Addi´ tionally, Curkovi´ c et al. [294] reported values of flexural strength in the range from 266.7 to 357.5 MPa for CIP high purity alumina ceramics sintered at 1650 °C. These results do not satisfy the values required by the ISO/DIS 6474-1:2019 standard [305] which specifies a flexural strength value of >500 MPa for ceramic materials based on high purity alumina for use as bone spacers, bone replacements, and components of orthopaedic joint prostheses. In spite of alumina ceramics having been continuously improved to fulfil clinical requirements, other ceramics are being strongly investigated, such as ZTA, and to a lesser extent ATZ composites and silicon nitride in order to reach load-bearing lifetimes of 30 years in vivo [306]. It is noteworthy that the mechanical properties, and more specifically, the flexural strength, of orthopaedic implants are strongly dependent on the starting materials and manufacturing methods. Therefore, within the scientific literature related to

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this topic, the results obtained are very varied and as such the in vitro experimentation must be extremely rigorous. As an alternative, zirconia ceramics which possess potential advantages in comparison to alumina in load bearing prostheses [307, 308], showing higher values in fracture toughness as well as flexural strength [300, 301], were introduced in the 1980s in response to concerns about Al2 O3 femoral head fractures in THA [265, 309], making it one of the most fracture-resistant ceramics. However, the clinical outcomes with ZrO2 heads were unpredictable with the literature reporting in vivo fractures of ZrO2 femoral heads [310– 312]. Moreover, catastrophic wear with ZrO2 -on-ZrO2 and ZrO2 -on-Al2 O3 articulations has been reported [313–315] along with the association of this material [265, 316] with high THA revision rates of the ZrO2 femoral heads used in THA in the 1980s, which in turn led to the FDA withdrawal of Y-ZTP from the market in 2001 [265]. In fact, those femoral heads still implanted in the patients were reviewed, or retrieved, considering that each implant needs to have an identification number to trace its origin in case of failure. Prior to its withdrawal from the market, the flexural strength values of Y-TZP for femoral head applications ranged from 1000 to 1500 MPa, values in agreement with the required by the International standard ISO13356 [317]. Under this scenario and considering that additionally Y-ZTP has a serious limitation in that it tends to degrade in a moist atmosphere at temperatures between 150° and 400 °C (this topic will be discussed later), nowadays it is preferably to combine the advantages of the two-single materials (Al2 O3 and ZrO2 ) by producing a mixed ceramic. According to Rahaman et al. [265], due to the limitations of Y-TZP, perhaps the most promising application of ZrO2 as a bearing material in THA and TKA appears to be its use as a particulate reinforcing phase in Al2 O3 to form ZTA composites, or alternatively, a ZrO2 matrix reinforced with Al2 O3 particulates to form an ATZ composite. Generally, in ZTA, the ZrO2 particles are mostly tetragonal, even without a chemical stabiliser if sintered properly. However, improvements in strength and fracture toughness are observed when the particles have a t-phase structure [265, 318–321]. ZTA offers improvements in flexural strength (700–1000 MPa) when compared with the average values for singlephase α-Al2 O3 (300–500 MPa) [265, 300] due to stress induced by the t to m phase martensitic transformation toughening, and microcrack toughening [322–324] where the extent of toughening achieved by ZTA is dependent on parameters such as the particle size of Al2 O3 and ZrO2 , the volume fraction of ZrO2 retained in the metastable tetragonal phase, the relative distribution of Al2 O3 and ZrO2 particles in the matrix, as well as chemical stabilisation [323]. Likewise, according to investigation by Garvie [325], it is inferred that the monoclinic phase is retained at room temperature owing to the presence of the hard alumina matrix, and this phase can be responsible for a significant improvement in the flexural strength transformation toughening mechanism. Moazzam Hossen et al. [322] found that the results of flexural strength testing showed a linear increase when the zirconia content and the sintering temperature were increased. For alumina with additions of 16 wt. % zirconia, the flexural strength was found to be

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around 383 MPa for pure Al2 O3 and 626 MPa for ZTA, sintered at 1450 °C, meanwhile 425 and 672 MPa were the flexural strength values obtained for the same ZTA composite but sintered at 1550 °C. This suggests that the martensitic transformation (t → m) could be the predominant mechanism to increase the flexural strength of ZTA composites. On the other hand, as the tetragonal zirconia grains do not undergo that stress-induced transformation to the monoclinic form when grain size is smaller than a critical size, the value of flexural strength is not affected. Nevertheless, the effect of grain size on flexural strength in transformation-toughened composites is a very complex phenomenon. Moreover, the flexural strength is strongly dependent on the manufacturing quality as well as the indentation method used, according to Gottwik et al. [326] who used the indentation in bending method. Depending on the alumina content in the zirconia matrix, the properties of the starting materials, as well as the procedures carried out to manufacture the ZTA composites, there is a wide spectrum of flexural strength values emphasising that not all of these values are suitable for its use in THA. As mentioned in Sect. 6.4, with the aim of producing a material with good biocompatibility, excellent chemical and hydrothermal stability, and high resistance to wear as well as good mechanical features such as toughness and flexural strength, the BIOLOX®delta project emerged combining the resistance and toughness of zirconia with the better wear resistance and toughness of alumina, to create a composite material using an alumina matrix [327–330]. In order to achieve the mentioned properties, small amounts of additives such as SrO were added to the base ZTA composition in order to increase the fracture toughness [265, 331–333]. The small Cr2 O3 content is not functional but gives the material its characteristic pink colour. During high-temperature sintering, these additives react with the Al2 O3 matrix leading to the in-situ formation of a small fraction of plate-like Al2 O3 grains which are uniformly distributed [330]. It is thought that the toughness increase with platelets may be caused by higher grain boundary adhesion with Sr and/or a change of internal residual stresses [334]. In this context, by controlling the amount of ZrO2 , SrO, and the sintering conditions, an optimisation of the microstructure is achieved which leads to the production of a material with a flexural strength greater than 1200 MPa. There are long term clinical outcomes with BIOLOX®delta and arthroplasty registries showing that these have the best clinical outcomes [334]. A flexural strength value near to 1800 MPa has been reported in hipped ZTA for applications where higher mechanical strength is critical [335]. Alternatively, and although there are very few studies carried out on ATZ composites, it is important to stress that this one can be a new material with a high potential in joint replacement. Indeed, Begand el al. [287] reported flexural strengths calculated according to ISO 6474 [305] of 1093, 1022, 962, 893, and 422 MPa for non-aged ATZ composites aged at 180 °C 5 h, 134 °C 30 h, and 134 °C 100 h, respectively, in comparison to the reference Al2 O3 Bionit®. These results are promising ones for the manufacture of products that have so far been the domain of metallic materials alone, and where the high strength is required. It is worth mentioning that a wide range of possible mechanical

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properties in ZTA and ATZ composites, with specific flexural strength, is largely the result of manufacturing processes that involve the development of an integrated production process beginning with the conditioning of the raw materials through to the final component. Undoubtedly, sophisticated procedures are used to achieve a significant improvement in mechanical properties, quality, and production efficiency, as well as cost [336]. On the other hand, the non-oxide Si3N4 has been used as an orthopaedic biomaterial to develop bearings that could improve the wear as well as longevity of prosthetic hip and knee joints [337]. However, Si3 N4 has been implanted in human patients for over 3 years now but clinical trials in prosthetic hip replacements using femoral heads manufactured with this ceramic are barely contemplated, taking in to account that the flexural strength is very varied in the resulting material due to the different options of fabricating components with silicon nitride. For example, the silicon nitride components obtained by means of the reaction bonding process show relatively low density, high porosity (typically 15–20%) and a low flexural strength of 200–300 MPa [337]. Conversely, Si3 N4 with additions of Y2 O3 and Al2 O3 increase the flexural strength up to 923 ± 70 MPa [338]. Si3N4 composites, when thoroughly investigated in terms of their mechanical properties and suitability for in vivo implantation, may play a very important role in the biomedical field [339] in spite of being relatively expensive, but their price versus performance is better than the alternatives [340]. Nonetheless, Si3 N4 manufacturing is complex, and this increases its cost.

8.2.3

Fracture Toughness

It is well documented that the low fracture toughness of ceramics remains a concern for the reliability of ceramic bearings. Therefore, the improvements in materials are a challenge in order to further reduce the risk of ceramic-bearing failures in THA [265]. With regards to this topic, the materials science makes a fundamental distinction between fracture strength and fracture toughness. On the one hand, fracture strength is referred to as the maximum mechanical stress a material can withstand without fracturing. On the other hand, the fracture toughness is the resistance of a material to the propagation of cracks. For example, ceramic materials that have been in use for a number of years, such as BIOLOX®forte, already have very high fracture strength; meanwhile BIOLOX®delta also shows an extremely high fracture toughness. Therefore, and applying these differences, it has a much higher capacity than other ceramic materials to resist the onset of cracking and to arrest the propagation of cracks [334]. However, in the scientific literature the reports mention fracture toughness for the majority of ceramics used in THA and TKA. Nowadays, within the advanced ceramic field, sialons are unique materials comprising four constituent elements often alongside rare earth oxide sintering additives. Silicon (Si), aluminium (Al), oxygen (O), and nitrogen (N) are bound together in one of

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three distinct crystalline phases, exhibiting a broad range of high-performance thermomechanical properties [341]. As an alternative to these materials, there are Al2 O3 , ZrO2 and a combination of them (ZTA and ATZ). Superior fracture toughness becomes a necessary requirement for the selected ceramic in orthopaedic replacements mainly in THA and TKA, and this is especially true when joint micro-separation processes occur in overweight patients, according to Pulliam and Trousdale [342], as well as in Asian patients who have a particular problem where their squatting and setting attitudes require prosthesis with special designs and a high joint deflection, according to a report by Ha et al. [343]. The high fracture toughness must be retained for long lifetimes, such as expected in young patients [344]. The low fracture toughness of Al2 O3 (~4–5 MPa m1/2 ) has been a limitation in some orthopaedic applications. As a counterpart, the fracture toughness values for Y-TZP (6– 12 MPa m1/2 ) are approximately two to three times the values for Al2 O3 , making it one of the most fracture-resistant ceramics [265]. This improvement is reached with additions of 2 mol% Y2 O3 to obtain microstructures with a grain size of ~0.3 μm. An addition of 2 mol% Y2 O3 produces a higher fracture strength but with a decrease of fracture toughness [345]. In reference to pure alumina, the scientific literature has reported a variety of fracture toughness values. For example: Žmak et al. [346] reported fracture values ranging 5.23–6.36 MPa m1/2 for alumina produced by slip casting, Willmann [347] reported 4 MPa m1/2 for aluminas of the 1970s, 1980s and 1990s, Bocanegra-Bernal et al. [348] reported 3.4 MPa m1/2 for conventionally sintered alumina, Nevarez-Rascon et al. [349] obtained 4.2 MPa m1/2 in pressureless sintered alumina, and Bocanegra-Bernal et al. [278, 281, 283] reported 2.5 MPa m1/2 in alumina prepared by spark plasma sintering (SPS), 3.5 and 5.2 MPa m1/2 in pure alumina prepared by conventional sintering and sinter-HIP procedures, and 2.9, 2.4 and 3.0 MPa m1/2 in monolithic alumina prepared by conventional sintering, HIP and sinter-HIP routes, respectively. On the other hand, the reported fracture toughness values of 2 generations of medical grade alumina [350] are relatively low (3.0 MPa m1/2 and 3.2 MPa m1/2 for BIOLOX introduced in 1974 and BIOLOX®forte introduced in 1995, respectively) in comparison to the above reported values. As can be seen from the mentioned fracture toughness values, these can vary in great measure depending on the raw material as well as the manufacturing method. Another very important point to consider is that the fracture toughness of technical ceramics is nowadays measured by various methods like direct crack measurement (DCM), indentation strength in bending (ISB), and single-edge notched beam (SENB), amongst others. Nevertheless, in recent years the validity of the fracture toughness values reported in the literature are questionable when these are obtained by DCM, inasmuch as it is considered that the fracture toughness could be overestimated [351]. Regardless of the measurement technique and judging by the results of several researchers, the SENB method leads to a more consistent and accurate measurement of mechanical properties. However, independent of the criticism reported in the literature, undoubtedly the values thus obtained by DCM can be useful for qualitative comparisons between samples in the

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same study with an acceptable nanoscale dispersion [352], taking into account that all the models employed in the fracture toughness calculations are essentially based on the same parameters and the differences are related to equation coefficients [353]. Considering that the largest manufacturer of zirconia femoral heads recalled their products in 2001 due to the great rates of failure in vivo [311], this unfortunate event has opened a strong and controversial issue on the future of Y-TZP as a biomaterial and because of that, the use of Y-TZP has declined sharply. Although the fracture toughness values for zirconia reported in the literature usually range 5–12 MPa m1/2 , the above-mentioned failure events have led to some surgeons going back to other solutions [354]. In this context, contemporary trends in material development are focussed on improving the properties of existing materials as well as developing new ones to improve the performance of ceramics in THA [354]. The combination of alumina and zirconia to obtain advanced composites can be the ultimate choice to take advantage of the high hardness and wear resistance of alumina and the higher strength and fracture toughness of zirconia [354, 355]. Under this criterion, it is important to mention the large number of reports in the scientific literature revealing results of fracture toughness for ZTA composites where both the content of alumina and zirconia are varied leading to a wide range of values that are influenced, as mentioned above, by parameters such as processing methods of the composites, raw materials, as well as the technique employed to obtain the fracture toughness values. For example, in an interesting investigation carried out on ZTA composites, Žmak et al. [346] showed how the toughness of monolithic Al2 O3 ceramics can be improved by the addition of ZrO2 nanoparticles. Table 8.7 [346, 353, 356–362] shows the calculated fracture toughness values for different ZTA composites in comparison to monolithic Al2 O3 . From Table 8.7 it can be seen that independent of crack type and applied model to calculate the fracture toughness of ZTA composites prepared by the conventional sintering of green bodies, formed by the slip casting method, and by the slip casting forming method, there is an improvement in the toughness of alumina when zirconia particles are added to its matrix. This alumina matrix ensures high hardness of the ZTA composites, while the addition of zirconia particles promotes the resistance to crack propagation [363]. The resulting ZTA composite also slows down the kinetics of hydrothermal ageing (discussed later on) which is an excellent advantage over monolithic ZrO2 . It has been mentioned previously that additions of small amounts of MgO ( 22° indicating that high acetabular component anteversion is associated with squeaking and consequently a high wear rate. It is important to stress that the optimum acetabular position and the relationship between this and the wear is highly dependent on the patient’s range of movement, which could be different for each one. Due to the already mentioned issues presented by alumina and zirconia monoliths in orthopaedic applications, the best balance between hardness, toughness and hydrothermal resistance has been reached thanks to alumina–zirconia composites in order to improve the ceramic properties [467]. Indeed, in the last fifteen years, the combination of these mechanical properties has led to the commercialisation of [215] ZTA composites improving their properties as a consequence of well dispersed and isolated zirconia grains constrained in an alumina matrix. Previous works [468, 469], have experimentally studied the effect of shocks as a source of wear damage on ZTA hip joints, where shocks are associated with micro-separation, that is, a separation between the femoral head and cup that induces short and high contact stresses between those two components at heelstrike, as pointed out by Dennis et al. [470] and Lombardi et al. [471]. Based on the above mentioned, an interesting investigation was carried out by Clarke et al. [472] under ‘severe’ micro-separation test mode by means of a CoC simulator with 36 mm BIOLOX® implants including ZTA composites (BIOLOX®delta, ‘d’) compared to alumina as the historical control (BIOLOX®forte, ‘f’). For this, they experimented four possible combinations (ball-cup: dd, df, fd and ff) which were run simultaneously in a hip simulator to 5 million cycles duration (5Mc) with femoral heads of 36 mm diameter. According to Asif [473], “the micro-separation occurs during the swing phase when the load is minimal and could occur clinically due to different factors such as head offset deficiency, laxity of the joint, medialised cup, impingement or subluxation”. The wear rates for all 36 mm CoC combinations tested by Clarke et al. [472] were very low, even under the severe micro-separation test conditions. The ff pairings demonstrated an ‘average’ wear rate of 1.5 mm3 /Mc whereas clinical wear rates for contemporary alumina THA retrievals ranged from 0.1 to 3.6 mm3 /year, which is within the mid-clinical range for alumina implants. The trend in this investigation was ranked as ff ≫ (df ≥ fd) > dd, in other words, ff pairs had the highest THA wear rates and dd pairs had the lowest. Likewise, making comparison between 36 mm versus 28 mm head size, the results are clinically negligible [472]. On the other hand, the micro-separation simulation research generated stripe wear on all

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ceramic components (BIOLOX®forte, BIOLOX®delta). Perrichon et al. [467] reported that shocks under micro-separation release wear debris of various shapes and sizes through inter and intra-granular cracks affected the hardness of the implant as well as the Young’s modulus. Studies on ZTA composites have revealed that shocks are the origin of wear stripes on the femoral head surfaces, and that they have to be considered in order to reproduce the in vivo degradation. The successful results with BIOLOX®delta have allowed the production of larger sized femoral heads with a reduced dislocation rate due to the reduced impingement and improved stability [474]. The reader is referred to consult Ref. [473] for a more detailed study of the wear of ZTA composites. Regardless of the drastic slow down after the catastrophic failure of specific batches of Y-TZP heads in the early 2000s which led to the premature LTD [467], isolated studies on the wear characteristics and clinical performance of zirconia femoral heads articulating against UHMWPE [475] have been reported with some contradictory results, considering that, as previously mentioned, both zirconia against alumina and zirconia against zirconia present catastrophic rates of wear in vitro as reported by Willmann et al. [273]. For example, results by Stilling et al. [71] do not suggest any advantage of zirconia articulating against PE compared to CoCr heads, and moreover, they did not experience zirconia head fractures at midterm follow up. On the other hand, Hummer et al. [312] and Oldenburg et al. [476], reported cases of catastrophic results after revision of fractured zirconia ceramic heads. From many years ago, and more specifically in 1991, a new product called Hylamer was marketed by DePuy-Dupont Orthopaedics (Newark, New Jersey) which was expected to have better yield strength and less wear compared to UHMWPE, to be tested in articulation against zirconia femoral heads [477, 478]. The annual wear for zirconia articulating against UHMWPE was 0. 17 mm/year whereas the wear of zirconia vs Hylamer was 0.40 mm/year [475]. Considering the conflicting results reported about the performance of the zirconia femoral head articulating with both UHMWPE and Hylamer, it is suggested that these femoral heads should not be used in THA. Hylamer was withdrawn from the market after several studies demonstrated high rates of wear and revision, according to Livingston et al. [479] and Cohen [480]. It is well known that a great majority of failures in THA were strongly related to wear of the cup, head and liner, where the accumulation of wear particles can be linked to osteolysis which leads to aseptic implant loosening [481]. However, today with alumina matrix and Vitamin E-infused highly cross-linked polyethylene (VitE-XLPE), this is not an issue anymore because Rochcongar et al. [482] reported results confirming that wear rates were lower in HXLPE/VitE cups than in UHMWPE cups, suggesting that HXLPE/VitE cups may prevent osteolysis, implant loosening, and eventually revision surgery. However, a challenge would be to reduce the generation of wear particles from the implant surfaces and therefore, Si3 N4 , which has a low wear rate when sliding against itself, would probably be in the short term a possible candidate for THA, emphasising that wear particles of this non-oxide ceramic can be slowly dissolved in polar liquids with the potential to be resorbed in vivo so reducing the risk for aseptic loosening. According

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to Boshitskaya et al. [483], an advantage of Si3 N4 for applications in THA lies in that Si3 N4 -powders are dissolved in blood serum and gastric juices which suggests that the use of this ceramic material in THA may produce less wear particles which in turn would be biocompatible and biodegradable. Olofsson et al. [481] concluded that Si3 N4 sliding against Si3 N4 showed a low wear rate in bovine serum as well as phosphate buffered saline (PBS). In this context, this good material combination for THA is promising and possibly its performance could be much better as a coating on CoCr. Further investigations are needed with the purpose to know the true future of Si3 N4 as a ceramic material for applications in THA. Nonetheless, it is very important to point out that there is no need for new ceramic materials in THA taking into account that ceramics have shown to be successful in THA and therefore, it is preferable to expand their use to other applications and to abandon CoCr completely.

8.2.6

Accelerated Ageing Tests

Another problem with the use of zirconia lies in its spontaneous transformation to its stable monoclinic form under in vivo conditions, known as LTD [274, 292, 484, 485]. Known about for more than 30 years [486], this results in increased implant surface roughness, enhanced wear, weakening of the material, and eventual fracture [310, 487]. In simple terms, Y-TZP may perform better than Al2 O3 but it is very sensitive to industrial variations [292]. Thanks to excellent mechanical properties and tribological performance, ZrO2 was introduced as an alternative to alumina for the manufacture of femoral heads [488]. It is worth remembering that ZrO2 exists in three different forms of crystalline structure: (i) the stable and brittle cubic (c) form at 2370 ºC; (ii) the stable, monoclinic (m) form present at ambient temperature; and (iii) the hardest and strongest but metastable tetragonal (t) crystalline form [327, 467, 484, 489, 490]. Although zirconia can be stabilised with oxides such as yttria [491], ceria, magnesia or calcia [485], Y-TZP has been used extensively due to its fine grain size and superb mechanical properties. However, efforts in developing zirconia-based ceramics with improved reliability for application in THA seem unsuccessful inasmuch as concerns exist related to their phase stability, structural reliability, and long-term wear performance [492, 493]. It is important to note that under thermal or mechanical stress (stress-induced phase transformation), ZrO2 has the ability to transform from its metastable t phase at ambient temperature to its stable m phase, representing one remarkable (somewhat unexpected) finding in the ceramic field [494, 495]. In fact, in the 1970s, first Garvie et al. [496] and then Gupta et al. [497] showed that zirconia exhibits a transformation toughening mechanism that acts to resist crack propagation. The relevance here is that the volume increase related to the structural differences is the source of increased crack resistance propagation, partially closing the crack when propagating. From this discovery, extensive investigations have been developed to understand the underlying mechanisms and potential implications and applications of this

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important phenomenon [495]. Paradoxically, the mentioned transformation, which imparts high strength and toughness to the zirconia, can cause a large reduction in strength when this material is exposed to temperatures in the range of room temperature to 500 °C for an extended period of time, especially under humid conditions, which in turn promotes the grains to pop-up and pull-out, microcrack formation, as well as subsequent surface degradation [484, 495]. Nowadays, the exact mechanism for this LTD is reasonably well understood but some occasional reports of experimental results do not seem to support these theories. For example, given enough time, the degradation can take place at room temperature. It is was demonstrated that Y-TZP material containing a low (2.5 mol. 9%) yttria concentration presented spontaneous cracks in a large number of pieces when these were stored under office conditions for several years [498]. As a consequence, investigations of the rupture of hip joints prosthesis caused by the transformation have regained interest from a biomedical point of view, mainly for ZTA considering that today, pure zirconia is rarely used in THA. The LTD is time dependent [487, 499] in vitro and in vivo [500–503] occurring as a result of a slow process of t to m phase transformation of the grains on any surface in contact with water [504] or body fluids, through the nucleation and growth mechanism of grains [501]. Clarke et al. [416] argued the ageing kinetics can be sensitive to different microstructural factors. Moreover, parameters such as purity of material, yttria content, density, grain size as well as surface roughness are very important in maintaining the ZrO2 ceramic as Y-TZP. In fact, the hydrothermal decomposition reaction depends on the grain size together with distribution of the stabilising yttria within the zirconia grains [505]. On the other hand, Schubert and Frey [506] and Chevalier et al. [274] reported that LTD is induced by polar molecules (typically water) and as a consequence of that, the t phase is destabilised and transforms to m phase without any external mechanical stress. Besides, the volume expansion due to the transformed domains puts the surface of components under compressive stress leading to the opening up of grain boundaries, which in turn facilitates the interpenetration of fluid into the bulk of the material. This LTD phenomenon, along with in vivo failure of ceramic hip joints, has been the main reason for the downturn of Y-TZP for applications in THA [494, 507, 508]. Additionally, LTD is typically described by a nucleation and growth model according to Mehl, Avrami and Johnson [509, 510]. In spite of that LTD has not yet been fully understood and the reported results are not totally in accord with each other, but it is clear that the t to m phase transformation proceeds from the surface to the interior [486, 511–513]. Guilardi et al. [514] proved in their investigations that the t to m phase transformation is a time-dependent spontaneous mechanism which is accelerated in the presence of different stimuli (moisture, temperature, mechanical cycling), no affecting both strength and structural reliability of Y-TZP ceramic under different experimental regimens carried out in their investigations. Nonetheless, it must also must be considered that not all Y-TZP materials behave similarly when submitted to stimuli, insomuch as they present differences in terms of composition and

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grain size and therefore, these are prone to present a high resistance to the effects of LTD [514, 515]. According to Guo [508], LTD can be delayed by increasing dopant content, the reduction of grain size, as well as the creation of an inert surface layer that can be produced by a recrystallised tetragonal layer with finer grain size [516], or alternatively, a cubic surface layer [517]. Given around 20 years of clinical experience related to zirconia femoral heads as well as more than 600,000 heads implanted, the results about this topic are very varied, some successful in vivo [518], others with osteolysis and surface degradation [316, 494], but there are a few reporting degradation by ageing or LTD [354]. Nonetheless, most investigations on zirconia implants were processed when the ageing was not yet fully understood. Undoubtedly, different investigations have proven that the strength and fracture toughness of Y-TZP ceramics are dramatically reduced when they are used in humid environments at temperatures between 150 and 400 °C for long times [519–522]. In fact, Lange et al. [519] argued that Y2 O3 reacts with water to form Y(OH)3 which leads to the Y2 O3 not acting as a stabiliser so favouring the t to m phase transformation. Considering these arguments as well as the different typical degradation mechanisms suggested by several researchers, and reported by Guo [508] in zirconia ceramics, other materials must be investigated with the purpose to avoid the LTD. Indeed, ZTA ceramics have replaced previous material generations of monolithic alumina and zirconia in the ceramic field, as well as cobalt-chrome alloys in metals [435, 523, 524] for applications in orthopaedics. There are an important collection of results concerning investigations with ZTA composites under humid environments [279, 322, 336, 365, 367, 376, 467, 495]. For example, Bartolomé et al. [525] reported significant ageing (up to 40 vol% m-ZrO2 after 20 h of ageing treatment time) and microcracking at Al2 O3 grain boundaries for ZrO2 , higher than the percolation limit of 16 vol%, creating pathways for water diffusion from the surface towards the bulk. On the other hand, a homogeneous distribution of alumina grains act as a constraint to the zirconia grains, contributing to a high hydrothermal degradation resistance inasmuch as the t-ZrO2 phase can be retained. Reyes-Rojas et al. [526] investigated the ageing applied to Al2 O3 + 13 wt. % Y-TZP + 2 wt.% m-ZrO2 composite in steam at 134 °C under a pressure of 2 bar. They found that the fraction of monoclinic polymorph (obtained values by means of Rietveld analysis conducted on the θ–2θ XRD data) was of 2%, 2.3% and 2.6% for ageing times of 0, 40, and 125 h, respectively, compared to 84.8% for pure Y-TZP aged for 40 h. It is probable that with a homogeneous dispersion of ZrO2 into the matrix, the fine grain size, as well as the high Young’s modulus of the Al2 O3 matrix, can provide a constraining effect on the zirconia and could partially inhibit the phase transformation, which is in accordance with the report by Bartolomé et al. [525]. In other words, the mentioned homogeneous dispersion of zirconia into the alumina matrix facilitates the non-propagation of the transformation by the presence of Al2 O3 grains between zirconia ones. This composite, in terms of the monoclinic phase content, fulfils ISO 1336–2008 requirements for orthopaedic applications once their mechanical properties are valued.

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Perrichon et al. [467] studied the resulting morphologies and microstructures of ZTA composites (belonging to the latest generation of commercialised BIOLOX® delta implants, CeramTec GmbH (Plochingen, Germany)), after in vitro tests simulated the in vivo degradation. They focussed on the femoral head inasmuch as an analysis of the cup can require the development of very specific characterisation procedures due to its concavity, and visualised that the presence of both surface and sub-surface micro-cracks triggered the ZrO2 phase transformation in those worn areas, reducing further crack propagation which in turn could contribute to the long-term resistance of the ZTA composite material against hydrothermal ageing. However, the fact that a moderate increase of the monoclinic fraction after severe ageing tests is observed does not necessarily mean a decrease in mechanical properties for this specific type of composite [527], which today is the bio-ceramic gold standard for load-bearing components in THA [435]. This emphasises that the early scientific literature concerning ZTA ageing kinetics indicated that the t to m phase transformation was also sensitive to both composition and processing, according to Gutknecht et al. [528]. In spite of several predictions on LTD under 2 bar applied pressure at 134 °C, as well as controversial results, some independent research groups have subsequently reported that ZTA femoral head retrievals showed significantly higher amounts of polymorphic transformation [529–531] in comparison to reports by Guicciardi et al. [527], Gaillard et al. [532], and Muñoz-Tabares et al. [533]. Likewise, a strong increase of the polymorphic transformation of ZrO2 has been observed [435] in proximity of metal staining [534]. These authors suggested that the hydrogen generated by the reactions between metal and steam accelerated the flow of water molecules and free oxygen to vacancy sites in the zirconia lattice [535]. In spite of everything that happened, an extraordinary balance between SCG and LTD resistance [363, 490, 494] can be reached with ZTA composites which offer high mechanical properties and a complete stability for alternative applications such as knee prostheses or dental implants [531]. Nonetheless, Yildirim and Kern [536] investigated ZTA composites with a typical biomedical grade composition [537] and they revealed that these materials are very sensitive to changes such as stabiliser composition and sintering conditions, reaching great LTD resistance with a stabiliser concentration range between 1.1 and 1.2%. Indeed, almost entirely tetragonal materials with 1.2 and 1.4 mol% yttria are quite stable with a final monoclinic content < 20% for the longest exposure time, as can be seen in Fig. 8.11 [536]. Under this criterion, these ZTA composites can therefore be considered absolutely safe for THA applications and exposure times in vivo of several decades. It is worth noting that regardless of the published results on LTD of ZTA ceramics, the optimum range to obtain an important balance between good strength, fracture resistance, and sufficient LTD resistance, is very narrow. Given the excellent properties of ZTA ceramics and due to an increasing interest in aesthetics and concerns originated by toxic and allergic reactions to certain alloys, both patients and dentists have been looking for other alternatives. ZTA composites could be candidates for potential use in dental applications such as endodontic posts, crown and

8.2

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Fig. 8.11 Monoclinic fraction of 17ZTA (Al2 O3 + 17 vol% ZrO2 ) composite exposed to water vapour (134 °C/2 bar) at different exposure times and sintered at 1525 °C. From Ref. [536]. Yildirim N, Kern F. Mechanical properties and ageing resistance of slip cast and pressureless sintered ZTA— the influence of composition and heat treatment conditions. Sci Sint 2019;51:243–56. Published by the International Institute for the Science of Sintering under the terms and conditions of the Creative Commons—Attribution 4.0 International licence. Copyright © 2018 Authors

bridge restorations and implant abutments [538]. Last, but not least, the dental market is still led by titanium, and ceramics still have to gain the attention of the majority of dentists. In fact, Aragón-Duarte et al. [539] reported a monoclinic zirconia phase content of 7% in ZTA ceramics for dental applications aged under a solution of artificial saliva at 134 °C and 2 bar pressure, results which are promising for these applications. Some interesting results related this field are reported, for example, in [490, 540, 541]. As was mentioned in Sect. 6.4, a great challenge lies in being able to avoid LTD of the zirconia-based composites without compromising their mechanical properties. In this context, Estili et al. [542] have reported a content of 55 and 10% of monoclinic ZrO2 in ATZ composites sintered at the same temperature in an air atmosphere, using alumina and graphite powder as a powder bed, respectively. The most interesting and surprising of these results is the formation in-situ of a surface alumina protective layer (Fig. 6.7) when it is sintered in a graphite powder bed, which significantly delays the LTD. The thickness of the protective alumina layer could be varied considering the final sintering time. It seems that carbonaceous products could inhibit the transformation of t to m phase since small addition of MWCNTs to the ATZ composite (ATZ + 0.01 wt. % MWCNTs) produced a 7% of monoclinic ZrO2 phase compared to 10% produced by the same ATZ composite free of MWCNTs, sintered under identical conditions (under graphite powder bed). More recently, Bocanegra-Bernal et al. [543] corroborated the insitu formation of the mentioned alumina layer surrounding ATZ samples, sintering them

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for longer times compared to previous investigations [542, 544] under the presence of graphitic compounds and a low oxygen atmosphere. The layer formation mechanism can be due to the reaction of CO with Al2 O3 during the heating stage of the composites, where the formed volatile species of aluminium in contact with the specimen surface, and owing to the presence of CO2 , are oxidised to the non-volatile Al2 O3 , thus forming the surface layer. Undoubtedly, the occurrence of this phenomenon improves the hydrothermal stability enhancing the possibility of using ATZ in THA as well as other applications, such as solid oxide fuel cells, and thermal barrier coatings. Therefore, this new and economical proposed route can be of vital importance to reach at constant sintering temperature and different times, the option to in-situ form a dense and hard alumina protective layer with varied thicknesses which could avoid the creation of a path for the water to penetrate down into the specimen. Nevertheless, the most used ceramic material for orthopaedic applications is still ZTA composite. In fact, Johannes and Schneider [545] have reported hydrothermal stability in the HIPed ZTA composite (ZrO2 + 10 wt.% Al2 O3 ) aged at 134 °C in water vapour under 2 bar pressure, obtaining a monoclinic zirconia phase content of approximately 1.5% after 192 h in the hydrothermal environment. They argued that this LTD resistance is due to the small grain size (143 ± 40 nm). The alumina addition stabilises the tetragonal phase and it could work as a barrier for the propagation of the phase transformation into the bulk [546, 547]. In spite of successful results, it is generally necessary to carry out investigations of the LTD of each newly developed zirconia-based ceramic [547]. It is noteworthy that in spite of the promising results with respect to ZTA compounds, there is no doubt that ATZ composites show high potential for joint replacement applications, since this composite combines the hardness and wear resistance of alumina as well as the fracture toughness and biaxial bending strength of zirconia. From this point of view, a typical surface microstructure for ATZ composite presents a predominant phase of zirconia surrounded by grains of alumina (See Fig. 6.6a). However, for the ATZ composite obtained by Bocanegra-Bernal et al. [543] in their recent investigations, its typical microstructural surface comprises an alumina protective layer which covers all the surface of the ATZ composite into an “alumina” monolith with the excellent properties of zirconia (See Fig. 6.6b), so avoiding the water penetration into the bulk of ATZ composite and improving the LTD. However, strong investigations are needed in order to evaluate the mechanical properties before and after ageing tests. Really, the report concerning the LTD of ZTA highlights the reduced roughening of ZTA compared to Y-TZP [489, 526, 548] so corroborating the few available studies on explanted ZTA components that specifically examine this point, according to Gremillard et al. [549]. On the other hand, the reported scientific literature on LTD of ATZ composites is even less extensive and therefore, papers detailing the LTD on explanted ATZ femoral heads are non-existent. Conversely, literature can be found that refers to flat samples that at no time perfectly reproduce the final surface characteristics of implanted components such as roughness and residual stresses [548]. Regardless of numerous differences reported regarding materials for applications

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in orthopaedics by scientists, industrialists and especially, orthopaedic surgeons, zirconiabased ceramics [354] are the most important of them since it was introduced 20 years ago to solve the brittleness, as well as the potential failure of implants [550]. However, for the short and middle term it seems that a combination of Al2 O3 and ZrO2 can lead to the production of an advanced composite optimising the manufacturing processes and raw materials in order to get the best properties. The study and understanding of the LTD, which is considered the Achilles heel of these alumina–zirconia composites in orthopaedic applications is of paramount importance for long-term performance of THA.

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Outlooks and Horizons in Materials and Technologies

It is well known that metals and their alloys, ceramics and polymers have been extensively used as biomaterials in joint replacements. Given the different atomic arrangement in these materials, there are a wide range of physical, mechanical, and chemical properties. However, although many of them have been manufactured in the industry, only some can be used in the orthopaedic field, and more specifically in THA, due to their bioactive, bio inert or biodegradable properties, as well as attractive mechanical properties such as strength, stiffness, fatigue life and so forth [1–4]. For the case of metals, these properties are strongly governed by their composition, microstructure, and phases and therefore, considering these different properties they could provide different applications in orthopaedics [2]. Regardless of whether it is a metal, ceramic or polymer, the main goal in THA is to return a patient to activities of daily living and a range of motion in the absence of pain as soon as possible and thus, in order to achieve that goal, a constant development of the material candidates to be used in joint replacement is essential [1]. According to Yoruc and Sener ¸ [4], Miura et al. [5] and Guo et al. [6], the new developments in artificial hip joints are essentially focussed on mechanical strength, biocompatibility, bioactivity and alternative materials that can impart better wear resistance and mechanical reliability [7]. It is important to understand that THA is formed by two components: on the one hand a cup type and on the other a long femoral type where the head of the femoral element fits inside the cup enabling the articulation of the joint. For many years, these two parts have been fabricated from metals, ceramics, polymers, and composites. The use of polymeric materials only results in a very fragile structure not suitable to meet the requirement of stress deformation responses in THR components [8]. On the other hand, although metals possess extraordinary mechanical properties, their biocompatibility is poorer inasmuch as the release of dangerous metal ions leads to detrimental consequences, like eventual failure, and in many cases to the removal of the implant. In view

© The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 A. Reyes Rojas et al., Performance of Metals and Ceramics in Total Hip Arthroplasty, Synthesis Lectures on Biomedical Engineering, https://doi.org/10.1007/978-3-031-25420-8_9

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of these unfortunate events, ceramic materials such as alumina and zirconia which usually have excellent biocompatibility but poor fracture toughness, being prone to be brittle, have emerged. However, a combination of these two monoliths have led to the fabrication of ZTA and ATZ composites where ZTA is the more established material for THA taking advantage of the combination of its biocompatibility, mechanical strength, and toughness [9]. In fact, the current challenge is focussed on several studies leading to obtain a perfect balance of properties, the result of an interdisciplinary research involving a great effort of material scientists, biomedical specialists, pathologists and clinicians, aimed at producing an implant to serve for longer times without rejection [8]. Nowadays, the scientific literature reveals that the main materials used in THAs are titanium, cobalt-chromium, PE, and ceramics, respectively [1]. Moreover, when we mention THA there are several options from which to choose the bearing surface, such as CoC, CoP, MoP, each with its advantages and drawbacks [10]. It is also clear that the use of MoM hip resurfacing implants should be made taking into account the pros and cons, considering relevant factors such as age, gender, body size, physical fitness and lifestyle, being that MoM pose the highest risk of provoking undesirable allergic reactions, and if MoM is the best fit for a patient’s needs, the surgery must be carried out by an expert orthopaedic surgeon to minimise the risks [11]. On the other hand, when ceramics are used, there is a very rare risk for fracture whereas metal bearings cannot fracture but are more prone to loosen and could require a second (revision) surgery. In another scenario, MoM releases tiny particles of metal into the joint which is absent with ceramics. However, in the event that ceramic debris can occur, the body does not seem to react to these wear particles like it does to the less biocompatible metal debris. The orthopaedic community is inclined to use metal bearings with polyethylene liners in older, less active individuals who are going to want this surgery to be their last. When it comes to younger and more active patients, the choice of the best bearing must be decided on a case by case basis where both surgeon and patient should consider all the pros and cons of the different articulating options (MoM resurfacing only, CoP, CoC, MoP) according to age, activity level, bone density and seriously comparing the relative risks and advantages offered by MoM and CoC options, the most commonly used nowadays. The use of MoP for younger patients is getting rare. The potential complications and every effort made to prevent any anticipated problems must also be considered [12, 13]. To ensure implant longevity in THA, an improvement in the manufacturing process, raw materials, and testing methods in vitro that can yield almost the same results as in vivo, undoubtedly lead towards the attainment of these goals. Thus, the scientific community will need to focus on not only further reducing abrasive wear, but also on reducing stress shielding by newer materials, as well as new designs. Although this is not an issue at all, the near future target, which is becoming a business need, will be substituting CoCr in other applications. Ceramics should win over alternative materials such as poly-ether-ether-ketone (PEEK). In closing, a wide and comprehensive knowledge of

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the characteristics, advantages and disadvantages of each bearing surface is of paramount importance for the surgeon’s group who routinely perform THA. Likewise, personnel with high experience will help in selecting the best coupling for each patient with the unique purpose of providing the best long-term survivorship of the prosthesis, emphasising once again that the best option according to the latest technological advances can be CoC in young and active patients [10].

References 1. Hu C-Y, Yoon T-R. Recent updates for biomaterials used in total hip arthroplasty. Biomater Res. 2018;22:33. https://doi.org/10.1186/s40824-018-0144-8. 2. Yadav S, Gangwar S. An overview on recent progresses and future perspective of biomaterials. IOP Conf Ser Mater Sci Eng. 2018;404:012013. 3. Navarro M, Michiardi A, Castaño O, Planell JA. Biomaterials in orthopaedics. J R Soc Interface. 2008;27:1137–58. 4. Yoruc ABH, Sener ¸ BC. Biomaterials: in a roadmap of biomedical engineers and milestones, Kara S, editor. Intechopen; 2012. p. 67–114. 5. Miura K, Yamada N, Hanada S, Jung TK, Itoi E. The bone tissue compatibility of a new Ti-NbSn alloy with a low Young’s modulus. Acta Biomater. 2011;7:2320–6. 6. Guo S, Bao ZZ, Meng QK, Hu L, Zhao XQ. A novel metastable Ti-25Nb-2Mo-4Sn alloy with high strength and low Young’s modulus. Metall Mater Trans A Phys Metall Mater Sci. 2012;43:3447–51. 7. Al-Hajjar M, Jennings LM, Begand S, Oberbach T, Delfosse D, Fisher J. Wear of novel ceramicon-ceramic bearings under adverse and clinically relevant hip simulator conditions. J Biomed Mater Res Part B Appl Biomater. 2013;101:1456–62. 8. Aherwar A, Singh MK, Patnaik A. Current and future biocompatibility aspects of biomaterials for hip prosthesis. AIMS Bioeng. 2015;3(1):23–43. https://doi.org/10.3934/bioeng.2016.1.23. 9. Zinger O, Anselme K, Denzer A, Habersetzer P, Wieland M, Jeanfils J, et al. Time-dependent morphology and adhesion of osteoblastic cells on titanium model surfaces featuring scaleresolved topography. Biomaterials. 2004;25:2695–711. https://doi.org/10.1016/j.biomaterials. 2003.09.111. 10. Zagra L, Gallazzi E. Bearing surfaces in primary total hip arthroplasty. EOR. 2018;3:217–24. 11. https://copublications.greenfacts.org/en/metal-on-metal-implants/index.htm. Accessed 20 Feb 2020. 12. Milošev I, Kovaˇc S, Trebše R, Levašiˇc V, Pišot V. Comparison of ten-year survivorship o hip prostheses with use of conventional polyethylene metal-on-metal, or ceramic-on-ceramic bearings. J Bone Joint Surg. 2012;94(19):1756–63. 13. https://loptonline.com/patient-education/injuries-conditions/hip-research-articles/which-is-bet ter-metal-or-ceramic-bearings-in-hip-replacements-see-more-at-httpwww-midwestphysicalthe rapy-cominjuries-conditionshipresearch-articleswhich-is-better-metal-or-ceramic-bearings/. Accessed 20 Feb 2020.

Conclusions

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Joint replacement using different materials has been used over the past 50 years with outstanding success. Millions of people throughout the world have benefited from THA surgery, extending their mobility and quality-of-life by years. However, a clear conclusion can be the one reported by Kuncická et al. [1] who points out that “significant opportunities remain to improve joint replacements by improving the materials from which they are made. While metals, ceramics, and polymers each serve critical functions in implant systems, their respective roles and properties are shifting with the advancements in each category and the introduction of hybrid material systems”. After reviewing different metallic, ceramic, polymeric and composite materials used nowadays in THA, there remain issues that must be solved in order to ensure good pain relief, increasing more and more the activity levels in young patients undergoing hip replacement, as well as a longevity of the prosthesis, reaching a higher range of motion as well as a stability in those ranges. It was observed that the mechanical, material and processing issues are imperative in the design, selection and improvement in the fabrication of hip replacements. For the specific application such as THA, a correct choice of the specific material to fulfil the requirements of different standards for particular orthopaedic applications is necessary. The proper selection of metallic, ceramic, polymeric or composite material plays a crucial role for getting the combination of properties such as high strength, wear, and corrosion resistance, as well as biocompatibility. With the passage of time and balancing the pros and cons of the various materials used for applications in THA, it could be determined that the use of ceramic materials (alumina and ZTA composite mainly) is best suited to these types of applications. In closing, the innovations in the design and fabrication processes for the different materials for THA are raising the great reality of being able to obtain, in the medium term, implants with improved performance to match both the biocompatibility and mechanical complexity of the hip implants. However, to achieve this, surgeons in close alliance with biologists and engineers cannot be ignored and therefore, must © The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 A. Reyes Rojas et al., Performance of Metals and Ceramics in Total Hip Arthroplasty, Synthesis Lectures on Biomedical Engineering, https://doi.org/10.1007/978-3-031-25420-8_10

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be persuaded with the long-term durability and reliability of the available biomaterials. Everything mentioned leads in the future to see novel biomaterials being developed that will increase the lifespan of orthopaedic implants. Today it is not possible to fully ensure which material is definitely the one that will dominate orthopaedics.

Reference 1. Kuncická L, Kocich R, Lowe TC. Advances in metals and alloys for joint replacement. Prog Mater Sci. 2017;88:232–80. https://doi.org/10.1016/j.pmatsci.2017.04.002.