Investigation of Hafnium for Biomedical Applications

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LICENTIATE T H E SIS

Department of Engineering Scienceand Mathematics Division of Machine Elements Jorge Sin Investigation Rituerto of Hafnium for Biomedical Applications

ISSN: 1402-1757 Investigation of Hafnium ISBN 978-91-7439-680-5 (print) ISBN 978-91-7439-681-2 (pdf) for Biomedical Applications Luleå University of Technology 2013 Corrosion and Tribocorrosion in Simulated Body Fluids

Jorge Rituerto Sin

INVESTIGATION OF HAFNIUM FOR BIOMEDICAL APPLICATIONS

40 cps /eV 20

Hf Hf C Hf O Hf Hf Hf Hf Hf Hf 0 12345678910keV

Corrosion and Tribocorrosion in Simulated Body Fluids

JORGE RITUERTO SIN

Luleå University of Technology Department of Engineering Science and Mathematics Division of Machine Elements Cover ﬁgure: Schematic representation of an atom of hafnium. Title page ﬁgure: Energy dispersive X-Ray spectroscopy (EDX) spectrum of hafnium.

INVESTIGATION OF HAFNIUM FOR BIOMEDICAL APPLICATIONS Corrosion and Tribocorrosion in Simulated Body Fluids

http://www.ltu.se or by contacting Jorge Rituerto Sin,

[email protected]

This document may be freely distributed in its original form including the current author’s name. None of the content may be changed or excluded without the permission of the author.

Printed by Universitetstryckeriet, Luleå 2013

ISSN: 1402-1757 ISBNISBN: 978-91-7439-680-5 (print) ISBN 978-91-7439-681-2 (pdf) Luleå 2013 Luleå 2013 www.ltu.se Preface

The work presented in this Licentiate thesis has been carried out at the Division of Machine Elements at Luleå University of Technology (Luleå, Sweden) and at the Institute of Engineering Thermoﬂuids, Surfaces and Interfaces (iETSI) (Leeds, UK). This research has been funded by the European Regional Devel- opment Fund through the Centrum för medicinsk teknik och fysik (CMTF) and Kempe Stiftelserna. The Swedish Research School in Tribology is acknowledged for travel funding. I would like to thank my supervisors Associate Professor Nazanin Emami, for giving me the opportunity of working on this project, and Professor Anne Neville, for her valuable help and guidance. I would also like to thank Dr. Xinming Hu for his input to my work. Thanks go to Martin Lund and Tore Serrander for their assistance in the lab and Lars Frisk for valuable help with sample preparation. I am grateful to my friends and colleagues at Leeds University, especially James Hesketh for long discussions about tribocorrosion and for his help in the lab. I would also like to thank all my friends and colleagues at the Divi- sion of Machine Elements at Luleå University of Technology for providing a friendly and enjoyable place to work. I would especially like to express my gratitude to Gregory Simmons for his valuable feedback on my manuscripts and Silvia Suñer for all her support and advice that help me to improve every day. Finally, I would like to thank my family for endless support and encourage- ment.

Abstract

Metals have excellent properties, such as high strength, ductility and toughness, which make them the material of choice for many biomedical applications. However, the main drawback of metals is their general tendency to corrode, which is an important factor when they are used as biomaterials due to the corrosive nature of the human body. Titanium and titanium alloys are widely used in biomedical devices due to their excellent corrosion resistance and good biocompatibility. However, one of the disadvantages of titanium is its low wear resistance. Hafnium is a passive metal with a number of interesting properties, such as high ductility and strength, as well as resistance to corrosion and mechanical damage. Previous studies have shown that hafnium has good biocompatibility and osteogenesis. However, the behaviour of hafnium in biological environment has not been studied in great depth. Furthermore, little is known about the resistance of the passive layer under wear-corrosion conditions and the effect of proteins on its corrosion and tribocorrosion behaviour. The overall goal of this study is to assess the potential of hafnium for use in biomedical applications. The aim of this work is to investigate the corrosion resistance of hafnium in simulated body ﬂuids as well as its behaviour in wear corrosion and fretting corrosion conditions. The results showed that hafnium presents a passive state in the presence of proteins and its oxide layer provides high protection to corrosion. In addition, although the passive layer could be disrupted due to wear and fretting, increasing the corrosion of the metal, it was rapidly rebuilt when the damaging ceased. On the other hand, the main drawback of hafnium was its tendency to suffer from localised corrosion. Although the formation of corrosion pits was retarded in the presence of proteins, it was drastically increased when hafnium was scratched or subjected to fretting.

Appended Papers

[A] J. Rituerto Sin, X. Hu, N. Emami. Tribology, Corrosion and Tribocorrosion of Metal-on-Metal Implants. Published in Tri- bology - Materials, Surfaces and Interfaces, 2013, 7(1), 1-12∗. Presented at the International Conference on BioTribology, September 18-21, 2011, London, UK. ∗Reprinted with permission of Maney Publishing and Tribology - Ma- terials, Surfaces and Interfaces. Permission to reprint this material is gratefully acknowledged.

[B] J. Rituerto Sin, A. Neville, N. Emami. Corrosion and Tribocor- rosion of Hafnium in Simulated Body Fluids. Submitted for journal publication. Presented at the 9th World Biomaterials Congress, June 1-5, 2012, Cheng- du, China.

[C] J. Rituerto Sin, A. Neville, N. Emami. Fretting Corrosion of Hafnium in Simulated Body Fluids. To be submitted. Presented at the 3rd International Tribology Symposium of IFToMM, March 19-21, 2013, Luleå, Sweden.

Contents

I The Thesis 1

1 Introduction 3 1.1 The need for biomaterials ...... 3 1.1.1 Types of biomaterials ...... 4 1.2 Metallic biomaterials ...... 5 1.2.1 Titanium and titanium alloys ...... 7 1.2.2 Possibilities of hafnium ...... 8 1.3 Tribocorrosion of metallic implants ...... 10 1.3.1 Corrosion of implant materials ...... 10 1.3.2 Wear of implant materials ...... 11 1.3.3 Case study: Tribocorrosion of Metal on Metal implants 12

2 Objectives 15

3 Materials and Methods 17 3.1 Materials and material preparation ...... 17 3.2 Test conditions ...... 17 3.3 Experimental set up ...... 18 3.3.1 Electrochemical cell ...... 18 3.3.2 Tribocorrosion set up ...... 18 3.4 Experimental methods ...... 19 3.4.1 Electrochemical tests...... 19 3.4.2 Tribocorrosion tests ...... 20 3.4.3 Fretting corrosion tests ...... 21 3.5 Surface analysis ...... 22 3.5.1 Microhardness ...... 22 3.5.2 Optical microscopy ...... 22 3.5.3 3D optical proﬁlometry ...... 23

vii 4 Results and Discussion 25 4.1 Corrosion resistance of hafnium ...... 25 4.2 Tribocorrosion of hafnium ...... 28 4.3 Fretting corrosion of hafnium ...... 30

5 Conclusions 35

6 Future Work 37

References 39

II Appended Papers 43

A Tribocorrosion of MoM implants 45

B Corrosion and Tribocorrosion of Hafnium 79

C Fretting Corrosion of Hafnium 97 Part I

The Thesis

Chapter 1

Introduction

In this chapter, the need for biomaterials is introduced. The most commonly used biomaterials are summarised, giving especial attention to the main groups of metals used for biomedical applications. Titanium and titanium alloys are discussed in more detail, and the possibilities of hafnium as an implant material are explored. Finally, the environment and conditions that biomaterials face when implanted into the body are discussed.

1.1 The need for biomaterials

The human body is formed by a number of organs that work together as one unit. When one or more of these organs are damaged due to injuries, infections or different types of diseases, body function is affected. Different solutions have been developed to minimise the adverse effect of these injuries and diseases. Often, these solutions involve the implantation of biomedical devices in the patient’s body in order to replace or improve the function of human tissue.

Biomaterials can be deﬁned as materials used to make devices to replace a part or a function of the body in a safe, reliable, economic, and physiologically acceptable manner [1]. The implantation of biomaterials in the body implies different restrictions that must be considered to minimise the risk of failure and the need of revision of the implant (Figure 1.1). A biomedical device should not generate any adverse biological response after implantation. Therefore, biomaterials must be biocompatible, non toxic and non carcinogenic. Also, the possible degradation products generated during the lifetime of the implant must not lead to any adverse reactions. In addition, the physical and mechan-

3 4 CHAPTER 1. INTRODUCTION

Figure 1.1: The main causes of implant failure that lead to revision surgery often combine biological aspects with material and engineering aspects. ical properties of the material must be adequate for the requirements of each particular application. The machining properties, the cost and the availability are also factors to be considered.

1.1.1 Types of biomaterials Biomaterials can be classiﬁed in ﬁve different groups:

• Metals: Metals are strong, tough and ductile materials. Their properties can be normally improved by alloying and by thermal and mechanical treatments. However most of the metals are highly reactive and therefore they can corrode unless they are protected. Screws for dental implants and fracture ﬁxation devices are some of their applications.

• Ceramics: Ceramics are stiff, hard, biocompatible and resistant to corrosion. However, their brittleness can be a concern when used for some applications, such as load bearing implants. The crown of dental implants and the femoral head of certain types of hip implants can be made of ceramic materials.

• Polymers: Polymers are light, resilient and easy to shape. However, polymers might be susceptible to fatigue cracking and in vivo oxidation. 1.2. METALLIC BIOMATERIALS 5

Implantable ocular lenses and artiﬁcial vascular grafts are some of the areas of application.

• Composites: composites combine two or more materials to obtain a material with improved characteristics. Their properties can be adapted according to certain requirements, however they can be difﬁcult to produce and they are normally expensive. Composites are used for different applications such as spine cages and acetabular cups for hip implants.

• Natural biomaterials: Beside the groups of synthetic materials, materials derived from animals or plants can be grouped as natural biomaterials. The materials included in this group can be more similar to the body, they can provide a better integration and, normally, they are less toxic. On the other hand, they are usually complicated to manufacture. They can denature and decompose more easily than most synthetic materials and they may have problems with immunogenicity. One example of this group is collagen, which has been used in tissue engineering for applications such as skin replacement, bone substitutes, and artiﬁcial blood vessels and valves [2].

Today, there is no biomaterial that can be universally used in all applications. The requirements of each particular application must be studied in depth in order to select the best candidate. The behaviour of the candidate material under the speciﬁc conditions must be understood in order to improve the outcomes of the biomedical implants. The ﬁeld of biomaterials is highly interdisciplinary and requires collaboration between clinicians, biologists, materials scientists and engineers in order to develop new biomaterials and improved medical devices.

1.2 Metallic biomaterials

Metals are generally deﬁned as polycrystalline compounds formed by metallic bonding, where an electron cloud surrounds positively charged metal ions (Figure 1.2) [3]. Metals present excellent mechanical properties such as high strength, ductility and toughness, which make them the materials of choice for various biomedical applications. Among the multiple uses of metals, cranial plates, dental implants, fracture ﬁxation plates and screws, joint implants are common. Metallic biomaterials are also used for various devices such as cages for pumps, valves and heart pacemakers, where materials do not support high loads. Various types of surgical instruments are also made of metal. 6 CHAPTER 1. INTRODUCTION

Figure 1.2: Schematic representation of a metallic bonding, where the elec- trons (red) freely move around the nuclei (green).

In spite of their good mechanical properties, the main drawback of metals is their tendency to corrode. Most of the metals used for biomedical applications are passive metals, an oxide ﬁlm is spontaneously formed on their surface which protects the material from the corrosive media.

In the 1900s, a metal alloy known as "vanadium steel" was speciﬁcally designed to be used for temporary ﬁxation plates [4]. Over the years, vanadium steel was replaced with more resistant materials. There are three main groups of metals that are used in medical applications: stainless steels, cobalt based alloys and titanium based alloys.

• Stainless steels: Stainless steel was the ﬁrst metallic material success- fully used as an implant [5]. Today, among the different compositions, the most commonly used stainless steel for biomedical applications is the 316L. The presence of chromium provides the passive state to the alloy. However, stainless steel tends to corrode under stresses and in oxygen depleted regions. Therefore, it is susceptible to suffer from crevice and pitting corrosion [5]. For this reason, stainless steel is normally only used in temporary devices such as fracture ﬁxation plates and screws, as well as for surgical instruments. 1.2. METALLIC BIOMATERIALS 7

• Cobalt-based alloys. The low wear resistance of stainless steel led to the development of a cobalt-chromium-molybdenum alloy known as Vi- tallium [6]. Today there are two main categories of cobalt-based alloys, Cobalt-Nickel-Chromium-Molybdenum (CoNiCrMo) alloys and Cobalt-Chromium-Molybdenum (CoCrMo) alloys, however, CoCrMo alloys are the most commonly used. The presence of chromium as an alloying element gives these alloys excellent corrosion resistance, which, together with their good wear resistance and mechanical properties, makes them suitable for applications where the device is subject to high loads. Therefore, it is the material of choice in total hip implants, where the femoral head is commonly made of Co-Cr-Mo alloy.

• Titanium-based alloys. Titanium and titanium-based alloys have excellent corrosion properties and biocompatibility for static implant conditions. However, their main disadvantage is their poor wear resistance. In addition to comercially pure titanium, the most commonly used titanium alloy is titanium-6% aluminium-4% vanadium (Ti-6Al-4V) [7]. Titanium and titanium alloys are used for many biomedical applications such as bone screws, dental screws, pacemaker cases, femoral stems in hip implants, etc.

1.2.1 Titanium and titanium alloys

Titanium was ﬁrst discovered by William Gregor in 1791 in Cornwall, Eng- land, but it was not extensively used until 1932 when William Kroll designed a process to produce metallic titaniun in industrial scale [8]. In the 1940s [9] titanium was studied as a possible implant material. When Professor Per- Ingvar Brånemark discovered that bone was able to form a direct interface with titanium (which he named as osseointegration) [10], titanium started to be extensively used in dental and joint replacement applications. Today, more than 1.000 tonnes of titanium are implanted into patients in the form of biomedical devices in the world every year [11].

Titanium is a highly reactive metal, it rapidly reacts with water molecules and forms an oxide layer on the surface. The passive layer that is formed is non porous and very adhesive, and it shows a low solubility. The formation of this oxide ﬁlm isolates the bulk material from the media, providing the metal with excellent corrosion resistance. The thickness of the protective ﬁlm is from 1 to 5 nm and it is mainly composed of titanium oxide [12], which is formed 8 CHAPTER 1. INTRODUCTION

"# !

Figure 1.3: Schematic representation of the reactions of the ions released into the body and their toxicity. according to the following oxidation reaction (1.1).

+ − Ti+ 2H2O ↔ TiO2 + 4H + 4e (1.1)

The reactivity of the metal and the released ions inﬂuence its toxicity and the tissue response (Figure 1.3). Ions can bind with biomolecules forming complexes that can activate the immune system. More active ions will react more easily with hydroxyl radicals and anions forming oxides and salts, decreasing the possibility of reacting with biomolecules. Less active ions will stay in the system longer and, subsequently, they will have more opportunities to react with biomolecules and reveal toxicity [13]. Therefore, the high reactivity of titanium not only provides it with good corrosion resistance, but also with low toxicity.

In spite of the good corrosion properties of titanium, it provides relatively poor wear characteristics [7]. Although titanium shows a low toxicity, the release of particles has been related to the release of osteolitic cytokines that eventually lead to implant loosening [14, 15]. Furthermore, they can be spread into the body and accumulate in various organs such as the liver and kidneys [16].

1.2.2 Possibilities of hafnium Hafnium is a metal included in group 4 in the periodic table, together with titanium and zirconium. The metal was ﬁrst identiﬁed by Dirck Coster and 1.2. METALLIC BIOMATERIALS 9

Georges de Hevesy in Copenhagen in 1923 [17] and owes its name to ’Hafnia’, the Latin name for Copenhagen. Hafnium is always found associated with zirconium in ores. The main mineral where it is found is zircon, with a ratio Hf/Zr of about 2.5% [18]. Due to a number of interesting properties such as high ductility and strength, as well as resistance to corrosion and mechanical damage [19], it has attracted interest for a number of applications. For instance, it is used as a control material for nuclear reactors an as an alloying element in some superalloys used in aircrafts engines [20].

In 1984, Marcel Pourbaix included hafnium among the group of metals to be considered for surgical implants due to the passive state presented by the metal. However, due to the lack of information about its toxicity to the human body at that time it was discarded from the ﬁnal list of metals to be theoretically considered [21]. The passive layer found on hafnium’s surface in aqueous solutions is mainly formed of HfO2, according to the equation (1.2). This oxide ﬁlm is highly stable in neutral solutions [22].

+ − Hf+ 2H2O ↔ HfO2 + 4H + 4e (1.2)

More recently, the properties of hafnium as an implant material have been investigated [23, 24]. Matsuno et al. [23] implanted a number of refractory materials (titanium, hafnium, niobium, tantalum and rhenium) in subcutaneous tissue and femoral bone marrow of rats, and performed histological observation and elemental mapping in order to evaluate the biocompatibility and osteogenesis of such materials. The results of the in vivo study showed that hafnium had good biocompatibility and osteogenesis. A different study on the tissue response to hafnium [24] showed a similar response in soft and hard tissues in two different animal species (rat and rabbit), which suggests that hafnium might be an interesting material for biomedical applications.

Hafnium has also been investigated as an alloying element in titanium alloys. Different Ti-Hf binary alloys have been reported in the literature. Wrought Ti-Hf alloys have shown a gradual increase in yield and tensile strength with increased amount of hafnium, up to 40%, whereas the ductility decreased [25]. More recent studies also showed an increased yield and tensile strength as well as bulk hardness for cast Ti-Hf alloys [26]. The electrochemical behaviour of this type of alloys is similar to pure titanium, showing an excellent corrosion resistance [27]. Ti-Nb-Hf-Zr alloys have also been investigated [28]. These alloys have shown a low elastic modulus which is beneﬁcial in order to reduce 10 CHAPTER 1. INTRODUCTION the stress shielding effect and to enhance bone growth. It has also been shown that cold work can be used to decrease the elastic modulus of this type of alloy, reaching values close to the elastic modulus of cortical bone [29].

To date, the behaviour of pure hafnium in biological environments has not been studied in depth. Little is known about the resistance of the passive layer to wear-corrosion conditions and the effect of proteins on its corrosion and tribocorrosion behaviour. Thus, further studies of hafnium under biological conditions are needed in order to determine the suitability of this material for biomedical applications.

1.3 Tribocorrosion of metallic implants

1.3.1 Corrosion of implant materials As soon as a medical device is implanted into the body it comes in contact with body fluids. The chloride ions dissolved in body fluids make them a corrosive environment for metallic materials. In addition, different organic species such as amino acids and proteins are also found in biological fluids, which influence the corrosion of metals [30]. Changes in the pH can also affect corrosion. The pH of the body normally remains around 7.4 because body fluids are buffered solutions [31]. The implantation of a device into a hard tissue can lead to pH values of around 5.6 or in the case of infection, the pH can reach values of around 9 [21].

Corrosion can be deﬁned as the degradation of a material caused by the reaction with its surrounding environment. The corrosion phenomena in human ﬂuids is due to an electrochemical processes, which implies two reactions: the oxidation of a metal, or anodic reaction (Eq. 1.3); and the reduction of an oxidising agent, or cathodic reaction (Eq. 1.4, 1.5) [31].

• Anodic reaction:

M → Mn+ + ne− (1.3)

• Cathodic reaction in neutral or alkaline solution:

− − O2 + 2H2O + 4e → 4OH (1.4) 1.3. TRIBOCORROSION OF METALLIC IMPLANTS 11

• Cathodic reaction in acidic solution: + − 2H + 2e → H2 (1.5) There are different types of corrosion that can affect materials. Some of the main types of corrosion are:

• General corrosion, also called uniform corrosion, affects the entire surface of the material, leading to the dissolution of metallic ions in the media.

• Pitting corrosion is a type of localised corrosion that consists of the formation of small cavities at the surface of the metal. The oxidation reaction in the pit leads to lower pH and the migration of chloride ions to balance the charge of the metal ions. This creates a more aggressive environment locally and enhances the corrosion, penetrating at a high rate.

• Crevice corrosion is a localised type of corrosion that occurs when a metal surface is partially shielded from the environment, for example, in narrow openings, joints or deep cracks. The anodic reactions preferentially occur in the crevice due to the difference in aeration with the bulk environment. Crevice corrosion can be minimised by optimising a design or proper material selection.

• Galvanic corrosion can take place when dissimilar materials are in contact and exposed to a conductive solution. Their different potentials tend to increase the corrosion rate of the more active material.

• Fretting corrosion takes place at the interface of contacting surfaces with small relative motion between them. It leads to an enhanced corrosion caused by the disruption of the passive layer.

1.3.2 Wear of implant materials Implant materials are subjected not only to the corrosiveness of the body but also to variable and complex service conditions. The static and dynamic loads affecting medical devices greatly depend on the activity of the patient. In addition, the application of the device will ultimately determine the type of conditions that will affect its degradation. Although passive state of metallic biomaterials minimises the failure rate directly caused by corrosion, different types of wear can damage the oxide layer and lead to accelerated corrosion 12 CHAPTER 1. INTRODUCTION

Figure 1.4: Combinde effect of wear and corrosion in metal on metal hip implants. processes. The bearing surfaces of joint implants are a clear example of the occurrence of wear in medical devices.

Deterioration by small relative motion between solids, normally known as fretting, can also be found in medical devices such as fracture ﬁxation screws and plates, at the head/neck interface of modular hip implants or at the cement/stem interface of hip implants. Fretting leads to mechanical damage of the oxide layer and wear of the metallic surfaces. The oxidised metallic particles also contribute to accelerated wear. The interaction between the mechanical contact, the electrochemical processes and the surrounding biological environment determines the material loss and the generation of metallic particles.

1.3.3 Case study: Tribocorrosion of Metal on Metal implants

Metal on metal (MoM) hip replacements have been considered as an alternative to metal on polyethylene (MoP) implants, especially in the case of young 1.3. TRIBOCORROSION OF METALLIC IMPLANTS 13 and more active patients who require a safe and long term performance of the device. MoM implants (Figure 1.4) consist of a metallic stem that is inserted into the femur, a metallic femoral head and a metallic acetabular cup that ar- ticulates with the head. The excellent wear properties of cobalt chromium molybdenum alloys makes it the material of choice for the bearing surfaces of the implant. It has been reported that MoM implants generate approximately 100-fold less volumetric wear than MoP [32] and the particle size has been estimated in the range of 60 to 90 nm [33], which ideally avoids the host response produced in the case of MoP. However, the mechanical damage caused on the bearing surface due to wear increases the corrosion of the implants and the release of metallic ions. Higher ion levels have been related with various biological reactions such as hypersensitivity, pseudotumours and carcinogenesis [34, 35, 36], which eventually lead to failure of the implant. Recent studies have also shown the importance of metallic ions generated in the contact between the cement and the femoral stem [37] and on the head and neck contact of modular implants due to fretting corrosion [38]. Today, MoM implants are the subject of controversy due to the concerns related the biological response to ions released into the body of the patient.

Chapter 2

Objectives

The main goals of this research work are summarised in this chapter.

The main objective of this work is to explore hafnium as a candidate material for medical devices. Hafnium has shown good biocompatibility and osteogenesis, comparable to that observed for titanium. Therefore it has been suggested as an interesting material for biomedical applications [24].

However, metallic implant materials are exposed not only to the corrosiveness of body ﬂuids but also to mechanical loads that can lead to wear. Little is known about the corrosion resistance of hafnium in biological environments, or its behaviour when subjected to different types of mechanical damage, such as wear and fretting.

The speciﬁc objectives of this research work were:

• To understand and summarise the importance of tribocorrosion in the ﬁeld of biomaterials. The particular case of metal on metal implants was studied in detail.

• To assess the corrosion resistance of hafnium in biological environments.

• To investigate the effect of wear on the corrosion resistance of hafnium.

• To assess the behaviour of hafnium in fretting corrosion conditions in simulated body ﬂuids.

Chapter 3

Materials and Methods

The materials and experimental methods applied in this research work are summarised in this chapter. The experimental set up used for this work is described, and the electrochemical techniques employed are explained.

3.1 Materials and material preparation

Commercially pure (CP) titanium (grade 2) and hafnium rods were studied in this work. Plates of 6 mm thickness and 25 mm diameter were prepared from the initial rod. The surface of the specimens was mirror polished and two surface ﬁnishes were studied (polished surface and scratched surface).

3.2 Test conditions

The rates of the electrochemical reactions are temperature dependant, therefore tests were conducted at 37 ◦C , which is considered as normal body temperature. In order to simulate body ﬂuids, solution of 25% bovine calf serum was used. To isolate the effect of the organic species (amino acids, proteins, etc.) from the saline environment, isotonic saline solution consisting on 0.9% NaCl was also used as a test solution. In addition, tribocorrosion experiments were performed in 50% bovine calf serum solution to magnify the effect of proteins on wear.

17 18 CHAPTER 3. MATERIALS AND METHODS

Figure 3.1: Schematic representation of the three electrode electrochemical cell used.

3.3 Experimental set up

The main goal of corrosion testing is to study the corrosion parameters of implant materials and to predict their behaviour when implanted into the body. Furthermore, new materials developed for medical applications must be studied not only in terms of corrosion resistance but the effect of wear and fretting should also be considered since they will affect the performance of the material.

3.3.1 Electrochemical cell

In order to study the corrosion behaviour of CP titanium and hafnium, a three electrode electrochemical cell was used (Figure 3.1). The electrochemical cell consisted of a platinum wire, that was used as the counter electrode; a Ag/AgCl electrode (3 M KCl), used as the reference electrode (196 mV vs SCE); and the studied specimens, that acted as the working electrode. All electrochemical measurements were carried out using a potentiostat.

3.3.2 Tribocorrosion set up

In order to investigate the tribocorrosion performance of CP titanium and hafnium, a three electrode electrochemical cell was integrated with a reciprocating tribological tester (Figure 3.2). The conﬁguration of the electrochemical cell was the same as the previously detailed. 3.4. EXPERIMENTAL METHODS 19

Figure 3.2: Schematic representation of the tribocorrosion set up.

Tribocorrosion tests The ball-on-plate sliding wear tests were performed in a reciprocating wear tester that conforms to ASTM G133. Specimens were mounted on a polyoxymethylene holder and were rubbed against a silicon nitride ball of 12 mm in diameter. A normal load of 40 N was applied. The frequency was controlled at 1 Hz and the stroke length was 10 mm. The specimens were passivated in 30% HNO3 for 30 minutes before each test according to the standard ASTM F-86. A digital microbalance with 0.01 mg resolution was used to gravimetrically determine the mass of the specimens.

Fretting corrosion tests A ball-on-ﬂat reciprocating tribometer was used to perform the fretting tests. Specimens were mounted on a polyoxymethylene holder and were fretted against a silicon nitride ball of 12 mm in diameter. A normal load of 10 N was applied and the frequency was controlled at 1 Hz. A maximum stroke length of 120 µm was used.

3.4 Experimental methods

3.4.1 Electrochemical tests Cyclic polarisation tests Cyclic polarisation tests are commonly used to assess the susceptibility of a material to suffer from localised corrosion [31], which is determined by the 20 CHAPTER 3. MATERIALS AND METHODS potential at which the anodic current rapidly increases. In addition, the rate of metal dissolution over a range of potentials can be studied. In order to carry out the cyclic polarisation tests, specimens were mounted in the electrochemical cell after polishing. The tests consisted of the following steps: • The open circuit potential was measured for 3600 s to stabilise the material surface. • The specimen was cathodically polarised at a scan rate of 10 mV s−1 down to -1000 mV. • The specimen was anodically polarised at a scan rate of 2 mV s−1 from -1000 mV up to 3000 mV or until the current density limit (500 µA cm−2) was reached. • The potential scan rate was reversed and the specimen was cathodically polarised to its open circuit potential value.

3.4.2 Tribocorrosion tests Open circuit potential (OCP) tests At the OCP, the net current is zero because the anodic reaction rate is equal to the cathodic reaction rate. The OCP gives a qualitative indication of the state of a metal (active or passive) in a particular environment (Figure 3.3) [39]. The OCP can be monitored while the materials are subjected to mechanical damage. A drop in the OCP indicates a more active surface, which suggests that the passive layer is being damaged. The OCP was monitored during the three stages of the test: • Specimens were immersed in the solution for 3600 s to stabilise them. • The tribological test started and was run for 3600 s. • The tribological test was ended and measurement was performed for 3600 s to study the repassivation of the materials. The percentage of repassivation of the OCP after mechanical depassivation had ended was evaluated according to Eq. 3.1: E − E Repassivation (%)= t wear (3.1) Epassive − Ewear

Where Et is the OCP value at a time t; Ewear is the OCP value during wear before the mechanical damage ended and Epassive is the OCP value in static conditions before the wear test starts. 3.4. EXPERIMENTAL METHODS 21

Loading Unloading

Passive material Potential (V)

Active material

Time (s)

Figure 3.3: Typical OCP values for passive material and active material in tribological contact.

3.4.3 Fretting corrosion tests

Cyclic polarisation tests Cyclic polarisation scans were performed on hafnium under fretting conditions to investigate the effect of fretting on the resistance of the material to localised corrosion. The tests consisted of the following steps:

• The open circuit potential was measured for 1800 s to stabilise the material surface.

• Fretting was started and the OCP was measured for 1800 s to ensure a stable value.

• The specimen was cathodically polarised at a scan rate of 10 mV s−1 down to -1000 mV.

• The specimen was anodically polarised at a scan rate of 2 mV s−1 from -1000 mV up to 3000 mV or until the current density limit (800 µA cm−2) was reached.

• The potential scan rate was reversed and the specimen was cathodically polarised to its open circuit potential value.

• Fretting was stopped. 22 CHAPTER 3. MATERIALS AND METHODS

Open circuit potential (OCP) tests The effect of fretting on the OCP was studied. The OCP was monitored during three stages:

• Specimens were immersed in the solution for 3600 s to stabilise them.

• The fretting test started and maintained for 3600 s.

• The fretting test was ﬁnalised and the OCP was measured for 3600 s.

Potentiostatic tests During potentiostatic tests a constant potential in the passive regime of the metal is applied and the anodic current is monitored. The mechanical depassivation produced by fretting and its effect on the current at a given potential can be analysed. A passive potential of 0 V (vs Ag/AgCl) was applied. The anodic current was measured during the test, which consisted on the following steps:

• The sample was immersed in the solution and a potential of 0 V (vs Ag/AgCl) was applied. The anodic current was monitored for 3600 s to ensure stable values.

• Fretting was started and the anodic current was measured for 3600 s to study the effect of the mechanical damage on the anodic current.

• Fretting was stopped and the currents were monitored for 3600 s.

3.5 Surface analysis

3.5.1 Microhardness Hardness can be deﬁned as the residual deformation of a material due to the penetration of another body. The microhardness of CP titanium and hafnium was studied with a microhardness tester using a load of 4903 mN according to the ASTM E-384.

3.5.2 Optical microscopy Optical microscopy was used to study the material surface after corrosion testing. The formation of corrosion pits was analysed using an optical microscope. 3.5. SURFACE ANALYSIS 23

Figure 3.4: Typical section of a wear scar obtained using 3D optical proﬁlom- etry.

3.5.3 3D optical profilometry 3D optical profilometry is a non contacting technique used to perform measurements of the surface topography (Figure 3.4). A Wyko 1100NT optical interferometer was used to study the cross section profile of the wear tracks after tribocorrosion tests.

Chapter 4

Results and Discussion

This chapter summarises the ﬁndings of this work. The corrosion resistance of hafnium is evaluated and compared to that of CP titanium, and the effect of wear and fretting on the corrosion behaviour of the materials is also discussed.

4.1 Corrosion resistance of hafnium

In this work, the corrosion properties of CP titanium and hafnium were studied in 0.9% NaCl solution and 25% serum solution. Cyclic polarisation scans were performed to investigate the susceptibility to localised corrosion and the rate of anodic dissolution over a range of potentials. The cyclic polarisation curves of hafnium in 0.9% NaCl and 25% serum are shown in Figure 4.1. A rapid increase of the current density was observed in both solutions, which was due to the breakdown of the passive layer and the formation of corrosion pits on the surface. Pits were observed under optical microscopy in 0.9% NaCl and 25% serum solution (Figure 4.2 (a) and (b)). The breakdown potential (Eb)was higher in the serum solution. Furthermore, the repassivation potential (Er), at which current density reaches passive values after pits are formed, was also increased by the presence of organic species in the solution. This could be due to the anodic polarisation of the specimens during the tests, which would attract the negative charged proteins [30] as well as the chlorides dissolved in the media. Chloride ions are known to be a cause of pit initiation, therefore, the competition between proteins and chlorides to adsorb on the surface would retard the formation of the pits [40]. When the current limit (500 µAcm−1) was reached, the potential slope was reversed. However, the current density continued increasing when the potential

25 26 CHAPTER 4. RESULTS AND DISCUSSION

3.0 3.0 Hf - 0.9% NaCl Hf - 0.9% NaCl 2.5 2.5 Hf - 25% serum Hf - 25% serum 2.0 2.0

1.5 1.5

1.0 1.0

0.5 0.5

0.0 0.0 0.0E+00 5.0E-04 1.0E-03 1.5E-03 2.0E-03 2.5E-03 3.0E-03 0.00E+00 1.25E-04 2.50E-04 3.75E-04 5.00E-04 6.25E-04 -0.5 -0.5

Potential (V) vs Ag/AgCl Potential (V) vs -1.0 Ag/AgCl Potential (V) vs -1.0 Current Density (A cm-2) Current Density (A cm-2) (a) (b)

Figure 4.1: Cyclic polarisation scans of (a) polished hafnium and (b) scratched hafnium.

(a) (b)

Figure 4.2: Corrosion pits observed on the surface of hafnium after cyclic polarisation tests on the mirror polished surface in (a) 0.9% NaCl and (b) 25% serum solution and on the scratched surface in (c) 0.9% NaCl and (d) 25% serum solution. 4.1. CORROSION RESISTANCE OF HAFNIUM 27

! " ! # ! # "

Figure 4.3: Anodic polarisation curves of CP titanium and hafnium in 0.9% NaCl and 25% serum. was decreased, which was due to the propagation of the pits. The pit propagation curve was also limited by the presence of organic species, which suggested that proteins could act as a diffusion barrier limiting the corrosion of hafnium when pits have been formed. A drastic reduction of the Eb was observed when imperfections were found on the material surface (Figure 4.1(b)), indicating a high correlation between surface ﬁnish and the tendency to suffer from pitting corrosion. Pits were preferentially formed around the surface imperfections as seen in Figure 4.2 (c) and (d).

Anodic polarisation curves, shown in Figure 4.3, were extracted from the cyclic polarisation scans. Both CP titanium and hafnium showed a similar behaviour, that can be divided in four zones with reference to the corrosion potential Ecorr (potential where the net current ﬂow is zero): at potentials below the corrosion potential (Ecorr), the current density was dominated by the cathodic reaction; at potentials around Ecorr there was a transition from cathodic to anodic currents; when the potential was increased above the Ecorr, both materials showed a passive plateau; ﬁnally, when the potentials were further increased, the Eb was reached and the current rapidly increased. In the case of CP titanium, the last phase was not observed and the passive layer remained stable at high potentials. Hafnium showed slightly lower values of current density in the passive state before the breakdown potential was reached. 28 CHAPTER 4. RESULTS AND DISCUSSION

(a) (b)

Figure 4.4: Evolution of the OCP during the tribocorrosion tests of (a) hafnium and (b) CP titanium.

4.2 Tribocorrosion of hafnium

The evolution of the OCP during the tribocorrosion test is shown in Figure 4.4. The resulting curve showed three zones:

• From A to B, the specimen was immersed in the solution and a relatively stable OCP value was reached due to the stabilisation of the passive layer.

• When the tribological test was started at B, the OCP immediately dropped towards more active OCP values, thereafter a more stable value was reached indicating the equilibrium between the mechanical depassivation and the electrochemical repassivation.

• Finally, wear was stopped at C and the OCP rose towards more noble values indicating that the passive layer was being reformed.

CP titanium showed more noble values than hafnium under the studied conditions. Both materials showed the ability to repassivate and rebuild the oxide layer once the mechanical damage stopped. The grade of repassivation was evaluated according to the Eq. 3.1 (Chapter 3) and it is shown in Figure 4.5. The OCP values recovered more rapidly in the case of hafnium, showing a slightly faster repassivation than CP titanium. The volumetric loss measured was higher in the case of hafnium in 0.9% NaCl and 25% serum under the studied conditions (Figure 4.6), which was inﬂuenced by the higher elastic modulus of hafnium (138 MPa [41]) compared to CP titanium (102.7 MPa [31]). This affects the initial Hertzian contact pressure, which was around 850 4.2. TRIBOCORROSION OF HAFNIUM 29

(a) (b) (c)

Figure 4.5: Evolution of the OCP of hafnium and CP titanium after the mechanical damage stopped in (a) 0.9% NaCl, (b) 25% serum and (c) 50% serum solution. 30 CHAPTER 4. RESULTS AND DISCUSSION

0.007 Hafnium 0.006

) Titanium 3 0.005 (cm 0.004 Loss

0.003

0.002 Volumetric

0.001

0 0.9%NaCl 25%serum 50%serum

Figure 4.6: Volumetric loss of hafnium and CP titanium after the tribocorrosion test in 0.9% NaCl, 25% serum and 50% serum solutions.

0.1 0.1

0.05 0.05

0 0 00.511.522.533.544.55 00.511.522.533.544.55 -0.05 -0.05

-0.1 -0.1 mm mm -0.15 -0.15

-0.2 -0.2 Hf - 0,9% NaCl CP Ti - 0,9% NaCl -0.25 Hf - 25% serum -0.25 CP Ti - 25% serum Hf - 50% serum CP Ti - 50% serum -0.3 mm -0.3 mm (a) (b)

Figure 4.7: Cross section proﬁle of the wear scar after the tribocorrosion tests of (a) hafnium and (b) CP titanium.

MPa in the case of hafnium and 750 MPa in the case of CP titanium. In addition, the cross section of the wear scar was also deeper in the case of hafnium under the conditions of this study (Figure 4.7).

4.3 Fretting corrosion of hafnium

Microhardness measurements

Microhardness measurements of CP titanium and hafnium are summarised in Table 4.1. Hafnium showed higher hardness values than CP titanium. 4.3. FRETTING CORROSION OF HAFNIUM 31

Table 4.1: Microhardness (HV500g) of hafnium and CP titanium. HV500g Hf 213 ± 3 CP Ti 163 ± 6 ±: Standard deviation

1.00

0.75

0.50

0.25

0.00 -0.0002 0 0.0002 0.0004 0.0006 0.0008 0.001 0.0012

-0.25

Potential (V) vs Ag/AgCl Potential (V) vs -0.50 Hf - 0.9% NaCl -0.75 Hf - 25% serum -1.00 Current Density (A cm-2)

Figure 4.8: Cyclic polarisation test of hafnium under fretting conditions in (a) 0.9% NaCl and (b) 25% serum.

Cyclic polarisation tests Cyclic polarisation tests were performed on hafnium to investigate the effect of fretting on the susceptibility of the material to localised corrosion. The breakdown potential of hafnium was dramatically decreased compared to the breakdown potential in static conditions (Figure 4.8). The effect of the organic species contained in the serum solution was similar to that previously reported. Due to the competition between negatively charged proteins and chloride ions to adsorb on the anodically polarised surface, the breakdown potential was higher in the serum solution in comparison to saline solution [40].

Open circuit potential tests The effect of fretting on the passive state of CP titanium and hafnium was studied by monitoring the OCP during the fretting test. The test can be divided into three stages as shown in Figure 4.9.

• From A to B the specimen was immersed in the solution and an anodic shift was observed, which was related to the stabilisation of the sample in the solution and thickening of the passive ﬁlm. 32 CHAPTER 4. RESULTS AND DISCUSSION

(a) (b)

Figure 4.9: Effect of fretting on the open circuit potential of hafnium in (a) 0.9% NaCl and (b) 25% serum solution.

Table 4.2: Maximum drop (ΔEmax) and variation (ΔEavg) of the open circuit potential due to the mechanical depassivation caused by fretting. ΔEmax (mV) ΔEavg (mV) Hf 0.9% NaCl 79.2 ± 25.9 36.9 ± 1.6 25% serum 152.0 ± 36.3 39.9 ± 19.4 CP Ti 0.9% NaCl 227.3 ± 43.6 81.3 ± 5.5 25% serum 260.4 ± 22.0 171.9 ± 21.3 ±: Standard deviation

• At B fretting started and the OCP dropped towards more active values and reached a minimum value. After that the OCP recovered and sta- bilised around less active value, which suggested a balance between mechanical depassivation and electrochemical repassivation.

• Finally, with the end of fretting at point C, the OCP value increased towards more passive values, indicating repassivation of the materials.

In spite of the more noble OCP values of CP titanium in static condition, the drop in the depassivation peak (ΔEmax) was more pronounced in the case of CP titanium in both 0.9% NaCl and 25% serum solution, as shown in Table 4.2. Furthermore, the average drop (ΔEavg) in the OCP caused by fretting was also higher in the case of CP titanium, suggesting that the passive layer formed on titanium suffered more severe mechanical damage than the oxide layer of hafnium. 4.3. FRETTING CORROSION OF HAFNIUM 33

Potentiostatic tests The evolution of the current during the tribological tests under passive potential (0 V vs Ag/AgCl) is shown in Figure 4.10. The test can be divided into three zones:

• From A to B the specimens were immersed in the solution and low anodic currents were observed due to the passive state of the metals.

• The onset of the fretting at B produced the removal of the passive layer and a rapid increase of the current was observed. After the initial peak, the current decreased towards lower values and an equilibrium between mechanical depassivation and electrochemical repassivation was reached.

• Fretting was stopped at C and the initial passive current values were rapidly reached and maintained until the end of the test at D.

The maximum anodic current (Ipeak), corresponding to the peak observed with the onset of fretting, and the average anodic current (Iavg) under fretting are summarised in Table 4.3. A higher peak was observed in the case of CP titanium in both 0.9% NaCl and 25% serum solution, indicating a more severe initial depassivation caused by fretting. The average current observed for CP titanium was also higher than for hafnium in the studied conditions. In the case of hafnium in 0.9% NaCl, the anodic currents reached a relatively stable value and no important variations were observed during the test. By contrast, for hafnium in 25% serum and CP titanium in 0.9% NaCl and 25% serum, the current rose during the test. The lower hardness of CP titanium could lead to greater mechanical damage and, subsequently, a more severe depassivation. Furthermore, proteins could have a negative effect on the repassivation of the metals restricting the diffusion of oxygen to the surface [30], which could affect not only the corrosion of the material but also the wear induced corrosion, leading to increased wear-corrosion damage. 34 CHAPTER 4. RESULTS AND DISCUSSION

Table 4.3: Anodic current peak (Ipeak) and average anodic current (Iavg) of hafnium and CP titanium during fretting at a passive potential (0 V vs Ag/AgCl). Ipeak (µm) Iavg (µm) Hf 0.9% NaCl 6.01 ± 3.02 0.86 ± 0.25 25% serum 4.35 ± 2.15 1.03 ± 0.22 CP Ti 0.9% NaCl 14.42 ± 6.16 2.82 ± 0.55 25% serum 23.67 ± 6.35 3.57 ± 0.52 ±: Standard deviation

(a) (b)

Figure 4.10: Evolution of the current as a function of time at a potential of 0 V in (a) 0.9% NaCl and (b) 25% serum. Chapter 5

Conclusions

The main conclusions of this work are summarised in this chapter.

Hafnium is a metal with good biocompatibility and ostegenesis, which make it an interesting material for biomedical applications. This work has investigated the corrosion behaviour of hafnium in simulated body fluids and the effect of wear and corrosion on the material degradation: • Hafnium showed a passive state in simulated body fluids. The oxide film formed on its surface provided high protection to corrosion. • Hafnium showed a higher susceptibility to suffer from localised corrosion than CP titanium. The organic species retarded the formation of pits and limited their propagation. • Surface imperfections increased the susceptibility of the material to pitting corrosion. Fretting also reduced its resistance to localised corrosion. • The passive layer of hafnium was disrupted by mechanical loads, leading to higher corrosion. However, the oxide layer was rapidly reformed when the damaging action stopped. • Hafnium showed lower anodic currents than CP titanium when exposed to fretting in simulated body fluids. Hafnium has shown that the highly protective oxide film formed on its surface in simulated body fluids provides it with excellent corrosion resistance, even when the material is subject to mechanical damage. However, its main drawback is the susceptibility of the material to localised corrosion, which is largely increased on scratched surfaces and under fretting.

Chapter 6

Future Work

This chapter suggests future objectives that could be pursued as continuation of this work.

This work has provided insight into the corrosion and wear corrosion behaviour of hafnium in simulated body ﬂuids. However, if hafnium is to be implanted into the body for a particular application, more realistic studies with speciﬁc geometry and loads, similar to those faced in vivo, should be performed.

In addition, further research is needed in order to investigate the effect of hafnium as an alloying element in titanium based alloys. Previous studies have shown that Ti-Nb-Hf-Zr alloys can have an elastic modulus close to the elastic modulus of bone [28], which would be beneﬁcial in order to reduce the stress shielding effect and to enhance bone growth. The resistance of this type of alloy to wear corrosion and fretting corrosion conditions should be studied in order to determine their suitability for orthopaedic applications.

References

[1] L. L. Hench and E. Ethridge, Academic Press, Inc., 1972, 385.

[2] C. H. Lee, A. Singla and Y. Lee, International journal of pharmaceutics, 2001, 221(1), 1–22.

[3] M. Pourbaix: ‘Atlas of electrochemical equilibria in aqueous solutions’, 1974.

[4] P. Parida, A. Behera and S. C. Mishra, International Journal of Advances in Applied Sciences, 2012, 1(3), 31–35.

[5] J. L. Gilbert and S. A. Mali: ‘Degradation of Implant Materials’, 2012, 1–28.

[6] C. Elias, J. H. Lima, R. Valiev and M. Meyers, Journal of the Minerals, Metals and Materials Society, 2008, 60(3), 46–49.

[7] M. Geetha, A. Singh, R. Asokamani and A. Gogia, Progress in Materials Science, 2009, 54(3), 397–425.

[8] W. Kroll, Transactions of the Electrochemical Society, 1940, 78(1), 35– 47.

[9] R. Bothe, L. Beaton and H. Davenport, Surg. Gynecol. Obstet., 1940, 71, 598–602.

[10] J. Wolfaardt, G. Gehl, M. Farmand and G. Wilkes, Int. J. Oral Maxillofac. Surg., 2003, 32(2), 124–131.

[11] Implant-based dental reconstruction: World dental implant and bone graft market, Technical Report 3rd ed, Rockville, MD: Kalorama Infor- mation, Kalorama Information, 2009.

39 40 REFERENCES

[12] T. Hanawa: ‘Degradation of Implant Materials’, Springer, 2012, 57–78.

[13] T. Hanawa, Mater. Sci. Eng., C, 2004, 24(6-8), 745 – 752.

[14] N. Taki, J. M. Tatro, J. L. Nalepka, D. Togawa, V. M. Goldberg, C. M. Rimnac and E. M. Greenﬁeld, J. Orthop. Res., 2005, 23(2), 376–383.

[15] S. Suñer, J. Tipper and N. Emami, Tribol. - Mater. Surf. Interfaces, 2012, 6(2), 39–52.

[16] R. Urban, J. Jacobs, M. Tomlinson, J. Gavrilovic, J. Black and M. Peoc’h, J. Bone Joint Surg. Am., 2000, 82(4), 457–457.

[17] D. Coster and G. de Hevesy, Nature, 1923, 111, 79.

[18] R. Tricot, J. Nucl. Mater., 1992, 189(3), 277–288.

[19] E. Cerreta, C. Yablinsky, G. Gray, S. Vogel and D. Brown, Mater. Sci. Eng., A, 2007, 456(1), 243–251.

[20] O. Levy, G. L. Hart and S. Curtarolo, Acta Mater., 2010, 58(8), 2887– 2897.

[21] M. Pourbaix, Biomaterials, 1984, 5(3), 122–134.

[22] W. A. Badawy and F. M. Al-Kharaﬁ, J. Mater. Sci., 1999, 34, 2483–2491.

[23] H. Matsuno, A. Yokoyama, F. Watari, M. Uo and T. Kawasaki, Biomate- rials, 2001, 22(11), 1253–1262.

[24] S. Mohammadi, M. Esposito, M. Cucu, L. E. Ericson and P. Thomsen, J. Mater. Sci.: Mater. Med., 2001, 12(7), 603–611.

[25] A. Imgram, D. Williams and H. Ogden, J. Less-Common Met., 1962, 4(3), 217–225.

[26] H. Sato, M. Kikuchi, M. Komatsu, O. Okuno and T. Okabe, J. Biomed. Mater. Res., Part B, 2005, 72(2), 362–367.

[27] Z. Cai, M. Koike, H. Sato, M. Brezner, Q. Guo, M. Komatsu, O. Okuno and T. Okabe, Acta Biomater., 2005, 1(3), 353–356.

[28] M. González, J. Peña, J. Manero, M. Arciniegas and F. Gil, J. Mater. Eng. Perform., 2009, 18(5), 490–495. REFERENCES 41

[29] M. González, F. Gil, J. Manero and J. Peña, J. Mater. Eng. Perform., 2011, 20(4), 653–657.

[30] M. A. Khan, R. L. Williams and D. F. Williams, Biomaterials, 1999, 20(7), 631–637.

[31] R. J. Narayan: ‘ASM Handbook, Volume 23 - Materials for Medical De- vices’, ASM international, 2012.

[32] H. Anissian, A. Stark, A. Gustafson, V. Good and I. Clarke, Acta Orthop., 1999, 70(6), 578–582.

[33] P. Firkins, J. Tipper, M. Saadatzadeh, E. Ingham, M. Stone, R. Farrar and J. Fisher, Biomed. Mater. Eng., 2001, 11(2), 143–157.

[34] H. Willert, G. Buchhorn, A. Fayyazi, R. Flury, M. Windler, G. Köster and C. Lohmann, J. Bone Joint Surg. Am., 2005, 87(1), 28–36.

[35] H. Pandit, S. Glyn-Jones, P. McLardy-Smith, R. Gundle, D. Whitwell, C. Gibbons, S. Ostlere, N. Athanasou, H. Gill and D. Murray, J. Bone Joint Surg. Br, 2008, 90(7), 847–851.

[36] R. Tharani, F. Dorey and T. Schmalzried, J. Bone Joint Surg. Am., 2001, 83(5), 774–780.

[37] M. Bryant, R. Farrar, R. Freeman, K. Brummitt and A. Neville, Tribol. Int., 2013, in press.

[38] A. M. Kop and E. Swarts, J. Arthroplasty, 2009, 24(7), 1019–1023.

[39] Y. Yan, A. Neville and D. Dowson, Wear, 2007, 263(7), 1105–1111.

[40] C. V. Vidal and A. I. Muñoz, Corros. Sci., 2008, 50(7), 1954–1961.

[41] S. D. Cramer, B. S. C. Jr and C. Moosbrugger: ‘ASM handbook volume 13B: Corrosion: Materials’, ASM International Materials Park, 2005. 42 REFERENCES Part II

Appended Papers

Paper A

Tribology, Corrosion and Tribocorrosion of Metal on Metal implants

A.1. ABSTRACT 47

Tribology - Materials, Surfaces and Interfaces, 2013, 7(1), 1-12

Tribology, Corrosion and Tribocorrosion of Metal on Metal implants

J. Rituerto Sin1,X.Hu2 and N. Emami1 1Luleå University of Technology, Division of Machine Elements, Luleå, SE-971 87 Sweden 2Wood Group Integrity Management, Middlesex, TW18 1DT UK

A.1 Abstract

Metal on metal joint replacements are considered as an alternative to metal on polyethylene implants, especially in case of young patients who require a safe and long-term performance of the device. The reduction of wear particles is a key factor in order to improve the life time of the implant in the human body. Metals have excellent properties that may increase the long-term success of the artiﬁcial joint replacement. However, corrosion of the metallic implant leads to an increase of the ion levels in the body of the patient. Metallic ions may produce a host response that can induce a catastrophic failure of the implant. This review initially focuses on the consequences that the degradation of the metals used in orthopaedic implants have for the health of the patient, and the different biological reactions that lead to the failure of the implant. Parame- ters that affect the release of particles and ions into the body are discussed as well. Special ttention is given to the tribology, corrosion and tribocorrosion behaviour of metal on metal implants. Finally, an overview of mathematical models that have been used to model the behaviour of the implants are also presented.

Keywords: Total joint replacement, Metal on Metal, Metal ions, Biotri- bology, Corrosion, Tribocorrosion 48 PAPER A. TRIBOCORROSION OF MOM IMPLANTS

A.2 Introduction

Total hip replacements are considered to be one of the most successful surgical interventions of the 20th century due to their effective treatment of chronic pain and diseases such as osteoarthritis, rheumatoid arthritis, hip fractures and tumours. During surgery, the unhealthy hip joint is replaced with an implant in order to restore the function and activity of the articulation. Today, it is estimated that approximately one million artificial hip joints are implanted an- nually in the world [1]. Metal on metal (MoM) hip implants were widely used in the 1960s. In spite of the limitation in manufacturing techniques, some of the first generation implants were clinically successful for 20 years or more [2]. However, the incidence of failure due to loosening was high. Metal on metal models such as McKee-Farrar were eventually abandoned in favour of Charnley’s model. The Charnley low friction arthroplasty featured a metal on polymer bearing (MoP). It was conceived as a procedure for elderly patients of low activity levels, but, due to the high level of success and improvement in surgical methods, the indications for joint arthroplasty have been expanded to younger and more active patients [3]. The longevity of MoP implants in the case of young patients is decreased due to the increased activity of the patient. Polyethylene high wear rates provoke a macrophage reaction that results in osteolysis and the loosening of the implant. Revision of the implant is inevitable after a period of time, which is normally more expensive and complicated than the primary arthroplasty. Over the last two decades, interest interest has increased in the so called hard on hard bearings, which have been developed to improve the longevity of the implants and decrease the risk of revision. They present low wear rates, which could decrease the risk of a biological reaction that may lead to osteolysis. Ceramic on ceramic articulations are a promising alternative due to their good biocompatibility and wear resistance. Despite improvements in the mechanical properties of ceramics, structural fracture of ceramic implants remains a major disadvantage for this type of bearing material [4]. Furthermore, cases of audible squeak have been reported [5]. The use of different bearing materials (ceramic on metal) has shown reduced wear rate compared with metal on metal [6] and lower ion levels [7]. However, long term clinical studies are needed to support these promising results. Second generation metal on metal implants were manufactured with improved materials and design. Optimised manufacturing techniques allowed improved clearances and surface roughness. This type of implants have been implanted A.3. EFFECTS OF ION RELEASE INTO THE BODY 49 since the end of the 1980s [8]. Concerns about the release of metallic ions and nanoscale wear debris limit the use of this type of implants [9]. The topic has been treated not only in scientific literature but also in general articles [10].

A.3 Effects of ion release into the body

Some of the positive attributes of metal on metal implants are their excellent wear properties, the potential for improved stability and high range of motion with larger diameter femoral heads [11]. Furthermore, midterm results with current generation implants are promising. Metal on metal implants generate approximately 100-fold less volumetric wear than metal on polyethylene implants [12], the particle size has been estimated to be in the range of 60 to 90 nm, and due to the small size, they are distributed all around the body [13, 14]. The generation of metal degradation products results in elevations in patients’ blood and urine metal levels. This metal ion level increase is the result of the corrosion of the metallic surfaces but also by the dissolution of nano-sized wear debris. Nano-sized particles are highly reactive and can be corroded easily due to their high surface area to volume ratio [15]. Normally, metal on metal implants are made of cobalt-chrome-molybdenum alloys. Cobalt, chromium and molybdenum can be found in the body but they become toxic depending on the concentration [16, 17]. Safe in vivo concentration limits still undeﬁned, nevertheless there are some reference values established by international groups [18, 19]. However, it is thought that the ions-tolerance might be patient-speciﬁc.

A.3.1 Measurement of ion levels in the body

The levels of metal ions depend on the ion production, transport, and excretion, but the kinetics of these processes are not clear [20]. There are some different methods to measure the metal ion levels in the body, such as analysis of blood, urine, synovial fluid, and others. Blood and urine are commonly used due to the ease of specimen acquisition. The daily output of cobalt in the urine can be used as a surrogate marker of the implant wear because cobalt is the most abundant metal in the alloy, and it is more soluble than chromium which is retained in the local tissue [21]. How- ever, due to variations in hydration during the day, spot urine samples tend to yield more variable metal concentrations [22]. To date, one of the most extended methods to measure and compare metal ion 50 PAPER A. TRIBOCORROSION OF MOM IMPLANTS levels is ion concentration in serum. Strong correlation has been found between measured serum and joint fluid concentrations of metal ions, and also a significant correlation between the maximum wear-scar depth and the metal ion concentrations in both serum and joint fluid [23]. These correlations lead to the conclusion that serum ion levels of chromium and cobalt can be used to estimate the amount of wear taking place in metal on metal implants. A standard analysis method would be an interesting tool to compare the performance of different types of prostheses. The comparative information would be very valuable as long as parameters such as the patient demographics, activity level, renal function, and analytic technique are comparable. However, the link between ion levels, symptoms and joint performance is still unknown. Developing an understanding of this is going to be a key issue during coming years.

A.3.2 Effect of the bearing couple on the ion levels

Chromium, cobalt and molybdenum ion levels in patients with metal on metal implants have been largely compared with other bearing couplings and with persons without any implant. The analysis of metal ion levels of stable ceramic on ceramic implants shows a negligible ion release in this type of prosthesis [24]. Patients who receive metal on metal implants show signiﬁcanly higher metal ion levels than patients with metal on polyethylene and non-implanted persons [25, 26]. In the case of ceramic on metal prostheses the difference is lower, but the levels are still higher for metal on metal prostheses [7].

A.3.3 Effect of the head size on the ion levels

In general, the stability provided by larger femoral heads in total hip arthroplasty is better than smaller femoral heads [11]. Range of motion, head-to- neck ratios and jump distances are also larger. The amount of wear and corrosion is considered related to the area of the bearing surfaces [27] but a larger femoral head diameter would lead to improvements in lubrication conditions, which would decrease the wear [28]. This issue has not been fully resolved, while some authors determine that large- diameter bearings result in a greater exposure of cobalt and chromium ions than small- diameter bearings [29], other recent studies show that there is no difference in metal ion levels between patients with small- or large- diameter femoral heads [30]. This uncertainty is affecting the introduction of large-diameter metal on metal couples, as used in hip resurfacing. A.4. PROBLEMS RELATED TO METAL ON METAL IMPLANTS 51

Conservative metal on metal hip resurfacing arthroplasty has been demonstrated as an alternative in young and active patients. The retention of the upper femoral bone stock allows more options at a later time if future revision of the implant is needed. Furthermore, stress shielding in the proximal femoral shaft is avoided and the risk of dislocation is lower. It has been reported that the increase in Cr and Co concentrations in the case of large-diameter hip resurfacing arthroplasty is signiﬁcantly higher than the increase in a standard-diameter (28 mm) metal on metal total hip arthroplasty [31, 29] but other studies report no differences between them [27, 32, 33]. Moroni et al. [34] studied these differences at midterm and they reported no differences in ion levels between patients with either metal on metal Birmingham hip resurfacing or metal on metal total hip arthroplasty. The effect of the correct implantation must not be overlooked since it has a great impact on the implant’s performance, especially in the case of hard on hard prostheses. For instance, the abduction angle of the acetabular component affects the metal ion levels in hip resurfacing components [35]. A more steeply-inclined acetabular cup will lead to a higher risk of edge-loading.

A.4 Problems related to metal on metal implants

Particles generated by metal on polyethylene implants are within the size range required for phagocytosis by macrophages, which eventually leads to a process of aseptic loosening of the implant [36]. By contrast, particles generated by metal on metal implants are within the nano-metre range, which reduces macrophage reaction. However, these nanoparticles cause some other reactions that cannot be overlooked.

A.4.1 Cytotoxicity Particles and ions released by metal on metal implants are distributed through- out the body and can affect the viability and functionality of the cells. In vitro studies show that the biological response of human cells to metal particles depends on the particle size. Nanoparticles are more easily ingested by the cell than micro-sized particles and they are rapidly corroded, this leads to high concentrations of metal within the cell. This results in greater cell death than in the case of microparticles [37]. Cobalt nanoparticles present higher in-vitro cytotoxicity than chromium nanoparticles [38]. At the same time, the cytotoxic dose of cobalt nanoparticles is higher than the correlative cytotoxic dose of cobalt ions, but it is still not clear if the combination of the two results 52 PAPER A. TRIBOCORROSION OF MOM IMPLANTS the cytotoxic effects. The concentration range found in vivo in patients with pseudotumours is at least three orders of magnitude lower than the critical concentration found in vitro [39]. This ﬁnding suggests that prolonged exposure at lower concentrations may also affect the cells, which has also been observed in-vitro [1].

A.4.2 Hypersensitivity

Hypersensitivity is described as a cell-mediated lymphocyte-dominated immunological response [40, 41]. It involves mainly T-cell activation but also B-cell activation [42]. Some of the associated symptoms are idiopathic pain and aseptic loosening [43]. This has been clinically reported in some articles associated with both hip resurfacing and total hip replacement metal on metal implants [44, 45, 46, 47]. The frequency of this delayed cell-mediated response is still unknown but appears to be low, it has been estimated to be between 0.3% and 5% [48, 49] but it requires more in depth study because it leads to the failure of the implant. When this immunological reaction is detected, the implant must be replaced by another with different bearing materials [40]. The risk of surgical revision does not increase in patients with metal allergies [50], however, metal sensitivity rate in patients showing a peripros- thetic T-lymphocytic inﬂammation has been reported to be higher than in the general population [41, 51]. Methods to detect potential risk of immunological response have to be developed, but a postoperative metal sensitivity could indicate an immunological response to the implant. This reaction seems to be exclusive for metal on metal implants [52], however, it has been suggested that similar responses are also present, but less common, in other types of implants [43].

A.4.3 Pseudotumours

A pseudotumor is described by Pandit et al. [53] as a soft-tissue mass associated with the implant which is neither malignant nor infective in nature. Various symptoms have been have been associated with the presence of pseudotumours, such as pain, palpable lump, nerve palsy and dislocation. The estimated incidence is approximately 1% ﬁve years post-operation, but it could increase progressively with time. Glyn-Jones et al. [54] reported 4% cumula- tive revision rate at 8 years, being higher in women (6%) than in men (0.5%). The revision rate in patients with pseudotumours is at least 70%, and the out- come of these revisions is poor, requiring further revision in some cases [55]. A.4. PROBLEMS RELATED TO METAL ON METAL IMPLANTS 53

The cause of these pseudotumours is not fully understood. It could be caused by the excessive presence of wear particles due to some problem with the articulation but it could also be due to an allergic reaction to the presence of metallic particles. Reported cases occur mainly in the presence of metal on metal hip resurfacing [53, 56, 57, 58], it could be linked with the higher ion levels reported by some authors with this type of implants.

A.4.4 Osteolysis

Particulate wear debris produced by metal on polyethylene is responsible for the macrophage response that causes osteolysis and eventually long-term failure. It is due to the size range of the particles generated. The risk of osteolysis is lower in metal on metal implants because the nano-sized particles generated generally do not produce this biological reaction. Clinical studies on metal on metal implants have shown mostly good results but there are some cases of osteolysis [59, 60, 61]. The presence of osteolysis in metal on metal implanted patients has been associated with hypersensitivity responses to the particles and ions released by the prosthesis [62, 63, 64]. However it is not fully clear if the hypersensitivity leads to instability of the implant or, by contrary, the instability of the implant causes excessive wear debris which then leads to the macrophage reaction which causes the osteolysis [65].

A.4.5 Genetic effect

One of the main concerns in regards to metal on metal implants is the possibility of genetic effects. Chromium and cobalt present in metal on metal implants are potentially mutagenic metals [66, 67]. It has been shown that particles of cobalt chrome alloy can cause chromosome and DNA damage in human cells [68]. The damage caused by cobalt chrome alloy particles is significantly higher than other material particles such as alumina [1]. DNA damage was found when synovial fluid collected at revision surgery was cultured with human fibroblast [69]. Chromosome aberrations such as chromosomal translo- cations and aneuploidy increase in the peripheral blood of patients within two years after metal on metal joint arhroplasty [70]. However, similar increase is found in the peripheral blood of patients at revision surgery of metal on polyethylene implants [71] The long-term effects of these chromosome aberrations are still unknown, further studies are necessary to evaluate the risk and consequences of implanted metal on metal prostheses. 54 PAPER A. TRIBOCORROSION OF MOM IMPLANTS

A.4.6 Cancerogenesis

There are concerns that chromosomal aberrations and DNA damage caused by metal particles and ions may lead to carcinogenesis, but it is still not clear. On one hand, Tharani et al. [72] reviewed a number of different articles determin- ing that the relative risk of cancer in patients with hip or knee arthroplasty is not higher than the expected in the general population. Which is supported by more recent studies [73]. On the other hand, the risk of cancer in patients who undergo metal on metal implants was found to be higher than in patients with metal on polyethylene implants [74]. Kidney and prostate cancers, lym- phophoma or leukaemia have a statistically higher occurrence in patients with MoM implants [75, 76]. However, the risk of cancer could be higher for patients with rheumatoid arthritis or osteoarthritis [77] due to the inﬂammatory nature of these diseases. Understanding the long term systemic effects caused by the metal particles and ions released by metal on metal implants is very important for their future use.

A.5 Tribology of metal on metal implants

The study of lubrication, friction and wear in both natural and artificial joints in biological systems is known as biotribology. In natural joints such as the hip or knee, the bones are covered by cartilage and the articulation is lubricated by synovial fluid, which is confined by the synovial membrane (Figure A.1). Natural synovial joints are designed to perform under different condition over their lifetime, they have excellent characteristics such as low friction, shock absorption, and they provide high mobility and stability to the joint. In spite of the different lubrication modes present under more severe conditions, the deformation of articular surfaces and viscous effects of synovial fluid enhance the prevalence of elastohydrodynamic lubrication regime, preserving the low friction and low wear during walking [78]. Different types of diseases can affect the properties of synovial fluid as well as cartilage degradation. Pain can be produced due to direct contact between bone surfaces. The replacement of the damaged joint may be necessary in some cases. To date, cartilage cannot be artificially replaced, thus the tribological conditions in the joint vary greatly when a prosthesis is implanted. The fluid generated after the replacement, sometimes called pseudo-synovial fluid, shows similar properties to the original synovial fluid [79]. A.5. TRIBOLOGY OF METAL ON METAL IMPLANTS 55

Synovial fluid Cartilage

Femoral head Acetabulum

Ligaments

Synovial membrane

Figure A.1: Schematic natural hip joint.

A.5.1 Lubrication of Metal on Metal implants

In absence of cartilage, the lubrication regime determines the contact between the asperities of the bearing surfaces and, consequently the friction (Figure A.2) and wear produced on the articular surface. Typically, three different forms of lubrication are identified.In the boundary lubrication regime the load is carried by the contact between the asperities of the surfaces. In mixed lubrication, the order of magnitude of the film thickness is similar to the surface roughness, therefore the load is carried by both the asperities and the pressures generated within the lubricant. In the case of full film lubrication, the film thickness is larger than the surface roughness, there is no asperity contact so ideally there is no wear. Smith et al. [28] studied the lubrication regime depending on the head size (Figure A.3). The authors showed boundary lubrication regime dominated for smaller diameters (16 and 22.225 mm). In this case, reduction in head diameter results in a proportional reduction of wear. An increase in head size of few millimetres, to 28 mm head size, led to a substantial reduction in wear since the lubrication regime shifts to mixed lubrication. The larger femoral heads studied (36 mm) showed surface separation in some gait phases, which indicates fluid film lubrication. The corresponding wear rates were very low, 56 PAPER A. TRIBOCORROSION OF MOM IMPLANTS

Friction Coefficient, μ Boundary Mixed Hydrodynamic Velocity x Viscosity Pressure

Figure A.2: Stribeck curve. which conﬁrms the beneﬁcial tribological behaviour of larger femoral heads.

boundary mixed Volumetric wear

approaching fluid film Increasing diameter of femoral head

Figure A.3: The effect of femoral head diameter upon lubrication and wear in metal on metal total hip replacements [28]. Reproduced by permission of the authors.

A.5.2 Wear phases The wear of MoM implants presents two different phases, an initial running-in phase characterised by high wear followed by the steady-state phase where the wear rate decreases and reaches a more stable level (Figure A.4). This biphasic behaviour has been reported in a number of hip simulator studies [80, 81, 82] and conﬁrmed by some in vivo studies, where maximum levels are reached A.6. CORROSION OF BIOMEDICAL METALS 57

Running-in Steady-state Volumetric wear

Number of cycles

Figure A.4: Schematic wear phases of metal on metal implants. during the ﬁrst phase and they decrease until a more stable level is reached [83, 84, 85].

A.5.3 Wear mechanisms

The main source of wear debris in artificial joints is normally the articular surface, but it can be found in other parts such as the interface neck-head in case of modular implants. Typically, wear can be classified in three principal processes: abrasive wear, which can be due to either two or three body, adhesive wear and fatigue wear. Nevertheless other types of wear such as corrosive wear can be found, especially in presence of metallic surfaces. In vitro studies of the wear mechanisms present in metal on metal artificial joints have shown signs of abrasion, surface fatigue and tribochemical reactions but not adhesion [86]. These results have been validated through analysing the surface of retrieved implants [87].

A.6 Corrosion of biomedical metals

Corrosion plays a key role in the release of metal ions in biological environments. The metallic materials most commonly used for biomedical applications are stainless steel, cobalt chromium alloys and titanium alloys. The corrosion resistance of these materials relies on the passive layer formed on their surface in contact with corrosive environment. 58 PAPER A. TRIBOCORROSION OF MOM IMPLANTS

A.6.1 Passive ﬁlm composition The oxidation reaction of passive metals results in the formation of an oxide ﬁlm according to the reaction: + M + nH2O → MOn + 2nH + 2ne (A.1) The oxide layer formed on stainless steel and CoCr alloys is primarily composed by chromium oxide (Cr2O3). Titanium and titanium alloy layers are mainly composed of titanium oxide (TiO2) [88].

A.6.2 Passive film breakdown The resistance of the passive layer is important in terms of structural integrity and as well as for the release of degradation products. The susceptibility of the layer to breakdown depends on the difference between the resting and breakdown potential, the more similar these values are the more probable breakdown on small areas of the surface (pitting corrosion) will be. Stainless steels used for biomedical applications are likely to suffer from pitting corrosion because these two potentials present similar values. In the case of CoCr alloys the difference is higher but corrosion is still possible under some conditions. How- ever, localised corrosion is not common in CoCr alloys, but they typically fail by transpassive dissolution [89]. The change in the composition of the passive film in the transpassive region leads to an increased dissolution rate, which is mainly dominated by the dissolution of Co. The difference between titanium and titanium alloys’ potentials is very high so breakdown of the passive layer is not possible under biological conditions [90]. When load and motion are present on the surface, the passive layer may be mechanically removed and fretting- and wear-corrosion become more important. The behaviour of the released ions will influence their toxicity and therefore, the tissue response. Ions can bind with biomolecules forming complexes that can activate the immune system. More active ions will react more easily with hydroxyl radicals and anions forming oxides and salts, decreasing the possibility of reacting with biomolecules. Less active ions will stay in the system longer and therefore they will have more opportunities to react with biomolecules and reveal toxicity [88].

A.6.3 Galvanic corrosion Galvanic corrosion can be caused by the contact between different metals but also the contact between the same metal with parts of the surface under corro- A.6. CORROSION OF BIOMEDICAL METALS 59 sion and parts under tribocorrosion conditions. The contact between different metals results in a corrosive attack on the active metal and in relative protection of the more noble metal, which limits the application of some coatings on metallic materials. This type of galvanic contact is usually present in modular implants, for instance, in the neck-head contact. Under static conditions, a galvanic couple of cobalt alloy and titanium is stable and there is no accelerated corrosive attack. However, when stainless steel is combined with either cobalt alloy or titanium, the couple is unstable and more susceptible to corrosive attack [91]. In the case of a mismatch between the modular connection neck-head, relative motion may result in wear on the surfaces, which may affect the passive layer and thereafter the corrosive behaviour of the couple.

A.6.4 Effect of proteins on corrosion

A wide variety of proteins exist in the biological fluids where biomaterials are implanted, therefore the interaction between these proteins and implanted materials is an important aspect that should not be neglected. The effect of proteins on the corrosion depends on material due to the different affinities of proteins for different materials. Clark et al. [92] studied the effect of the proteins serum albumin and fibrogen on some pure materials. In the case of metals such as Co and Cu, an increased corrosion rate was reported. The effect on Cr and Ni was a moderate increase whereas the corrosion was inhibited with Mo. The effect on Al and Ti was little. Muñoz et al. [93] studied the effects of albumin and phosphate ions on the corrosion of CoCrMo alloy. The adsorption of albumin on the surface has a double effect, first as an inhibitor, impeding the access of the oxidant to the metallic surface. On the other hand, most of the proteins present a negative charge in the body’s pH 7.4, whereas the charge of metallic ions is positive. Proteins and ions can bind, increasing the dissolution rate [94]. At anodic potentials albumin accelerates corrosion but around corrosion potential the two effects compensate each other [93]. In the case of phosphate ions, its adsorption decreases the anodic reaction rate. Phosphate ions and albumin compete to adsorb on the surface, which will decrease the effect of phosphate ions ac- celerating the anodic reaction (Figure A.5). The effect of proteins on titanium is similar to the effect shown on CoCrMo before, the proteins can accelerate the dissolution of the metal but at the same time they adsorb onto the surface blocking the cathodic current. The affinity to adsorption of fibrogen onto ti- tanim surfaces is more than that of albumin [95], but the albumin layer seems to be more efficient at blocking the cathodic reaction [96]. In the case of tita- 60 PAPER A. TRIBOCORROSION OF MOM IMPLANTS nium alloys, proteins can enhance corrosion resistance of Ti6Al4V, but reduce the corrosion resistance of other alloys such as Ti13Nb13Zr and Ti6Al7Nb [97]. The dissolution of metals in presence of proteins is also reported for stainless steel. The dissolution of the nickel and molybdenum is increased by the presence of fibrogen. Transferrin and albumin increase the dissolution of chromium [98].

Figure A.5: Albumin inhibits cathodic reaction (arrow 1) but can accelerate the anodic oxidation by binding metal ions (arrow 2). Phosphate ions inhibit the anodic reaction (arrow 0) [93]. Reproduced by permission of ECS - The Electrochemical Society.

Equation A.2 has been used to describe the total material loss produced in tribocorrosion systems [99] T = W +C + S (A.2) where W is the material loss due to wear in absence of corrosion, C’ is the material loss due to corrosion in absence of wear, and S represents the synergy which includes the effect of wear on corrosion and the effect of corrosion on wear.

A.6.5 Tribocorrosion measurements Tribological tests are commonly combined with electrochemical experiments in order to study the tribocorrosion. Three-electrode cells are normally integrated in different types of tribometers to monitor the electrochemical behaviour under tribological contacts. A potentiostat controls and monitors the A.6. CORROSION OF BIOMEDICAL METALS 61

M+ Metal nanoparticles M+ Metallic ions + + M M+ M

Passivated Mechanical Electrochemical metal Depassivation Repassivation

Figure A.6: Schematic representation of the depassivation/repassivation process. electrochemical signals of the electrodes while the load and speed are controlled by the tribometer.

Tribometer

Simple reciprocating wear testers such as a pin on plate tester are commonly used [100, 99] to perform tribocorrosion testing. Micro-abrasion tests [101, 102] have been used to study the importance of two and three body abrasion in tribocorrosion systems. A pendulum friction simulator [103, 104] has been also used. The main difference between this conﬁguration and simpler reciprocating testers is that the pendulum friction simulator operates in the mixed lubrication regime instead of the boundary lubrication regime. More realistic hip simulators have been developed to perform in-situ electro-chemical measurements [105, 106].

Electrochemical cell

Electrochemical techniques are used to understand the effect of corrosion in the complex system. The open circuit potential gives limited information on the kinetics of surface reactions [107]. However, information about the corrosion state of the material (active or passive) can be obtained, as well as information about how the tribological contact affects the passivity of the material and how the passive layer is recovered once wear ceases. The potential shift can be used as an indicator of the damage of the passive layer (Figure A.7). Anodic Polarisation provides information about the rate of dissolution in a given environment [99]. The breakdown potential indicates the resistance of the passive layer to localized corrosion initiation. However, the non-stationary conditions of the test limit the physical interpretation and quantiﬁcation of the results. In potentiostatic tests the current is measured at ﬁxed potential as a function of time. The mechanical depassivation and the electrochemical repassivation can be evaluated following the charge transfer rate. Cathodic protection tests are 62 PAPER A. TRIBOCORROSION OF MOM IMPLANTS

Loading Unloading

Passive material Potential (V)

Active material

Time (s)

Figure A.7: Typical open circuit potential for passive material and active material under tribological contact. performed to isolate the effect of wear in the equation A.2 by completely stop- ping the degradation due to corrosion. The difference in material loss with and without application of cathodic protection can be assumed to be the corrosion- related damage [108].

A.6.6 Effect of proteins on the tribocorrosion behaviour As soon as the prostheses are implanted into the body, they are in contact with the proteins present in biological fluids, which will affect the performance of the device. The formation of a film derived from proteins has been reported. The proteins that play a role in the formation of this tribofilm vary, initially more abundant proteins dominate the process but they are replaced over time with proteins which are more affine for the implant surface [109]. While the exact bonding mechanisms are not fully understood, the presence of high number of free ions close to the surface could ease the formation of metal/protein complexes [110]. The stable layer acts as a solid lubricant between the bearing surfaces which produces beneficial effects for the wear behaviour, for instance the prevention of adhesive wear by avoiding direct metal-metal contact [86, 110]. The tribofilm acts as an effective charge transport barrier that limits the ion release [101, 106]. On the other hand, an increase in the dissolution rate [93] in the presence of proteins may enhance the corrosion-related wear [101, 99]. The characteristics of this film were reported by Yan et al. [106], who studied it in both in vitro and in vivo samples. The comparison between them showed A.7. MODELLING 63 similar characteristics on the macro-level but also in the thickness and composition. It was described as an organometallic film which contained Co and Cr. The thickness of the film was about 15-80 nm and it was similar in both samples. Similar results have been reported using different test configurations [87]. Recently, signs of a graphitic layer have been found on metal on metal bearing surfaces. However, the mechanisms of formation of such material are not clear [111]. The formation of the so called tribofilm should be further studied, the mechanical properties should be determined and how it influences the tribology of the bearing must be clarified.

A.7 Modelling

A.7.1 Lubrication models The lubrication regime that dominates in the articulating surfaces of MoM hip implants has a great inﬂuence on the amount of degradation products released into the body. If full ﬁlm lubrication is achieved, the direct contact between surfaces is minimized, which potentially reduces wear debris. The lubrication of MoM implants has been widely investigated using computational simulating techniques, that are a powerful tool for understanding problems and predicting results. Mathematical models have been developed to study the tribological performance of the bearing in MoM prostheses.

Elastohydrodynamic lubricaion

NOTATION cd radial clearance Rc cup radius e eccentricity Rh head radius E Young modulus t time f calculated load u entraining velocity h ﬁlm thickness w applied load hmin minimum ﬁlm thickness δ elastic deformation p pressure η viscosity R equivalent radius = 1/Rh − 1/Rc

Due to the geometry of the femoral head-cup contact, spherical coordinates are normally used to establish the main equations governing the EHL problems (Figure A.8) [112]. 64 PAPER A. TRIBOCORROSION OF MOM IMPLANTS

Figure A.8: Spherical coordinates.

The pressure of the fluid film in the bearing is governed by the Reynolds equation (Equation A.3). ∂ ∂ ∂ ∂ ( 3 p)+ θ ( 3 θ p)= ∂ϕ h ∂ϕ sin ∂θ h sin ∂θ ⎡ ⎤ −w (sinϕsin θ∂h + cosϕcos θ ∂h ⎢ x ∂θ ∂ϕ ⎥ η 2 θ + ( ϕ θ∂h − ϕ θ ∂h 6 Rc sin ⎣ wy cos sin ∂θ sin cos ∂ϕ ⎦ (A.3) ∂h +wz sinθ∂ϕ ∂h + 12ηR2 sin2 θ c ∂t The film thickness is usually expressed as shown in Equation A.4, where both rigid and elastic deformation between the two bearing surfaces are considered [112].

h(ϕ,θ)=cd − ex sinθcos ϕ − ey sinθsin ϕ − ez cosθ + δ(ϕ,θ) (A.4)

The elastic deformation of the surfaces is modelled by Equation A.5, where K denotes the influence coefficient of the elastic surfaces and θm denotes a fixed mean latitude [112]. δ(ϕ,θ)= K(ϕ − ϕ ,θ − θ ,θm)p(ϕ ,θ )dθdϕ (A.5) ϕ θ

The external load components ( fx,y,z) are balanced by the internal hydrody- A.7. MODELLING 65

Figure A.9: Geometry of ball-in-socket (left) and ball on plate (right) models.

namic pressures (px,y,z) (Equation A.6). = 2 θ ϕ = fx,y,z Rc px,y,zd d wx,y,z ⎧ ϕ θ ⎫ 2 ⎨px = psin θcosϕ⎬ (A.6) = 2 θ ϕ ⎩ py psin sin ⎭ pz = psin θcos θ

The two different models primarily used to perform elastohydrodynamic analyses of metal on metal hip joint replacements are the ball-in-socket and the ball on plate (Figure A.9). Although the geometric similarity with a real hip joint is higher in the ball-in-socket lubrication model, it has been demonstrated the capability of both to perform steady-state EHL analysis [113, 114]. The main difference between them was found for large cup inclination angles, where the ball-in-socket model predicted no change in film thickness while the ball on plane model predicted reduction in film thickness due to the lubricant supply at the cup edge. The approximation of synovial fluid with Newtonian, iso- viscous and incompressible fluid is normally used because it does not show significantly different results under typical operating conditions in comparison with more sophisticated non-Newtonian model that considers the shear thin- ning effect of synovial fluid [113]. Simulated operating conditions influence the results of the lubrication model. Elastohydrodynamic lubrication analysis has been performed under transient conditions, such as a walking cycle or the cycling gait [115, 116, 117]. Results obtained simulating the whole walking cycle show different film thickness in comparison with steady-state conditions [112], which shows that the results are highly influenced by the loading and motion conditions simulated and highlights the importance of realistic parameters. 66 PAPER A. TRIBOCORROSION OF MOM IMPLANTS

Mixed lubrication A model of mixed lubrication for metal on metal hip implants was presented by Wang et al. [118]. The governing equations were based on the EHL problem but the boundary contact of absorbed protein layer was also considered. Boundary lubrication was assumed to occur when the predicted ﬁlm thickness was less than the protein layer thickness, which was assumed to be 30 nm. Furthermore, the work analyses theoretically the friction and frictional torque present on the implant bearing surfaces, which allows for comparison between experimental and simulated values and, therefore a direct validation of the model. In spite of the importance of roughness on friction and lubrication, the authors assumed smooth bearing surfaces as well as constant thickness of the protein layer. As said in Section A.5, different studies have shown the possible existence of mixed lubrication in metal on metal hip implants, however, few studies have been presented in recent years due to the complexity of mixed lubrication modelling.

A.7.2 Wear prediction Computational methods have been also used to simulate the long-term effect on wear for metal on metal hip implants and to understand the parameters that may affect them during their life [119]. Wear models are normally reproduced as 3D ball-in-socket geometry. The model is normally governed by Archard’s wear law (Equation A.7), which calculates the volumetric wear (V) depending on the resultant normal load (W), the relative surface sliding distance (L), and a coefficient of wear (Kw). V = KwWL (A.7) Despite the incomplete description of wear mechanisms of Archard’s wear law, it has been used due to its simplicity and validity in wide applications. The models are dependent on the wear coefficient, which is usually obtained experimentally from wear testing using a pin on plate or hip simulator to adapt the model to the characteristics of the studied bearing. The variation of the wear coefficient between the running-in and steady state phase must be care- fully considered. The relationship between metal on metal volumetric wear and elastohydrodynamic film thickness was studied by Dowson [119], based on a number of hip simulator studies performed in different laboratories. The Hamrock- Dowson equation (Equation A.8) was used to derive the Equation A.9, the minimum film thickness depending on the diametral clearance and the head A.7. MODELLING 67 diameter. h ηu 0.65 w −0.21 H = min = 2.8 (A.8) min R ER ER2 2.19 = . · 8 d hmin 1 6904 10 0.77 (A.9) cd The experimental wear rates obtained from a number of simulator studies are plotted against the theoretical film thickness in Figure A.10. It is shown that the steady state wear rate is almost independent of the film thickness whereas the running in depends greatly on the film thickness. Low running-in and steady state wear rates can be observed for a predicted film thickness about 20-25 nm [119]. The volumetric wear produced during running in is equivalent to many years of operation of the implant in the steady state phase.

A.7.3 Tribocorrosion models As mentioned earlier, the total material loss is enhanced in corrosive media, however, the complexity of the wear-corrosion processes makes it difficult to model the parameters that control the problem. A number of different ap- proaches have been used by different authors. Mischler et al. [120] proposed a model that describes the anodic corrosion current (Ir) of passive metals sliding against a rigid non-conductive body as function of experimental parameters such as frequency ( f ), wear track length (L), load (W ) and hardness (H), as well as a proportionality coefficient Kr that cannot be theoretically determined(Eq. A.10). 1/ f 0.5 −0.5 Ir = Kr LfW H idt (A.10) 0 This model considers the contact between asperities, which is described as a function of load and hardness, to determine the depassivated area. The prediction of the effect of mechanical and physical parameters shows a good correlation correlation with the experimental results, however, the wear component was not modelled in this work. A model based on the concepts of micro- cracking and fatigue to model the wear related damage, that considers both corrosion, wear and the synergism between them was presented by Jiang and co-authors [121]. However, although several tribocorrosion models have been published in the literature, there is a need of bio-tribocorrosion models, where the kinetics and mechanisms that affect the process are influenced by the biological environment. 68 PAPER A. TRIBOCORROSION OF MOM IMPLANTS (a) ) 3 Running-in wear (mm

Elasto-hydrodynamic film thickness (nm)

(b)

Figure A.10: Steady state (a) and running-in (b) wear rates depending on the theoretical ﬁlm thickness [119]. Reproduced by permission of the authors. A.8. CONCLUSIONS 69

A.8 Conclusions

Metal on metal implants reduce the volume of wear debris and the size of the particles released into the body, which ideally avoids the host response produced in the case of metal on polyethylene implants and, consequently, osteolysis and loosening of the implant. The longevity of the implant can be theoretically improved. However, the advantages of metal on metal implants may turn into disadvantages. The levels of metallic ions are increased in the body and these ions do not remain in the area close to the implant but they are dispersed around the body.

• While in some patients the high levels of metallic ions do not have any consequence, in some other patients it can lead to different biological reactions that can cause pain and failure of the implant. The causes of these types of host reactions are not fully understood. Prediction of which patients are more likely to suffer from severe biological reactions should be investigated deeply.

• The decrease of degradation products would minimize the host response, further research in the degradation mechanisms of the implants is needed.

• Tribology and corrosion are key factors in terms of release of ions and wear particles. The synergism between them in biological environments needs to be better understood and studied.

• Improved surgical techniques would further improve the outcomes of hard on hard implants, reducing cases of enhanced degradation due to malposition of the prosthesis.

• New materials with improved tribological and corrosion properties which lead to decreased degradation products and improved outcomes are de- sirable.

In spite of the concerns related with metal on metal hip implants, they still are a potential alternative for total joint replacements, therefore, more ef- forts should be made to understand and predict their behaviour and biological response. Achieving this requires strong collaboration between a variety of scientiﬁc ﬁelds and researchers. 70 REFERENCES

A.9 Acknowledgements

The author would like to thank Gregory Simmons of the Division of Ma- chine Elements at Luleå University of Technology for valuable feedback on the manuscript.

References

[1] A. Tsaousi, E. Jones and C. Case, Mutat. Res., 2010, 697(1), 1–9.

[2] S. Jacobsson, K. Djerf and O. Wahlström, Clin. Orthop., 1996, 329, S60–S68.

[3] S. Kurtz, E. Lau, K. Ong, K. Zhao, M. Kelly and K. Bozic, Clin. Or- thop., 2009, 467(10), 2606–2612.

[4] Y. Cai and S. Yan, Orthop. Surg., 2010, 2(3), 175–181.

[5] C. Jarrett, A. Ranawat, M. Bruzzone, Y. Blum, J. Rodriguez and C. Ranawat, J. Bone Joint Surg. Am., 2009, 91(6), 1344–1349.

[6] P. Firkins, J. Tipper, E. Ingham, M. Stone, R. Farrar and J. Fisher, J. Biomech., 2001, 34(10), 1291–1298.

[7] G. Isaac, C. Brockett, A. Breckon, D. Van Der Jagt, S. Williams, C. Hardaker, J. Fisher and A. Schepers, J. Bone Joint Surg. Br, 2009, 91(9), 1134–1141.

[8] C. Delaunay, I. Petit, I. Learmonth, P. Oger and P. Vendittoli, Orthop. Traumatol. Surg. Res., 2010, 96(8), 894–904.

[9] R. Crawford, C. Ranawat and R. Rothman, J. Arthroplasty, 2010, 25(1), 1–2.

[10] B. Meier, New York Times, 2010.

[11] A. Lombardi, M. Skeels, K. Berend, J. Adams and O. Franchi, Clin. Orthop., 2011, 469(6), 1547–1553.

[12] H. Anissian, A. Stark, A. Gustafson, V. Good and I. Clarke, Acta Or- thop., 1999, 70(6), 578–582. REFERENCES 71

[13] P. Firkins, J. Tipper, M. Saadatzadeh, E. Ingham, M. Stone, R. Farrar and J. Fisher, Biomed. Mater. Eng., 2001, 11(2), 143–157.

[14] R. Urban, J. Jacobs, M. Tomlinson, J. Gavrilovic, J. Black and M. Peoc’h, J. Bone Joint Surg. Am., 2000, 82(4), 457–457.

[15] G. Mabilleau, Y. Kwon, H. Pandit, D. Murray and A. Sabokbar, Acta Orthop., 2008, 79(6), 734–747.

[16] T. Rae, J. Bone Joint Surg. Br, 1981, 63(3), 435–440.

[17] A. Vyskocilˇ and C. Viau, J. Appl. Toxicol., 1999, 19(3), 185–192.

[18] D. Henschler, G. Lehnert and D. F. K. zur Prüfung Gesundheitss- chädlicher Arbeitsstoffe: ‘Biological exposure values for occupational toxicants and carcinogens: critical data evaluation for BAT and EKA values’, V. 1, VCH, 1996.

[19] M. Morgan, Environ. Health Perspect., 1997, 105(Suppl 1), 105–115.

[20] C. Heisel, M. Silva, A. Skipor, J. Jacobs and T. Schmalzried, J. Bone Joint Surg. Am., 2005, 87(4), 781–787.

[21] J. Daniel, H. Ziaee, C. Pradhan, P. Pynsent and D. McMinn, J. Bone Joint Surg. Am., 2010, 92(4), 840–845.

[22] S. MacDonald, W. Brodner and J. Jacobs, J. Arthroplasty, 2004, 19(8), 12–16.

[23] K. De Smet, R. De Haan, A. Calistri, P. Campbell, E. Ebramzadeh, C. Pattyn and H. Gill, J. Bone Joint Surg. Am., 2008, 90(Supplement 4), 202–208.

[24] L. Savarino, M. Greco, E. Cenni, L. Cavasinni, R. Rotini, N. Baldini and A. Giunti, J. Bone Joint Surg. Br, 2006, 88(4), 472–476.

[25] L. Savarino, D. Granchi, G. Ciapetti, E. Cenni, A. Nardi Pantoli, R. Ro- tini, C. Veronesi, N. Baldini and A. Giunti, J. Biomed. Mater. Res., 2002, 63(5), 467–474.

[26] V. Rasquinha, C. Ranawat, J. Weiskopf, J. Rodriguez, A. Skipor and J. Jacobs, J. Arthroplasty, 2006, 21(6), 47–52. 72 REFERENCES

[27] A. Moroni, L. Savarino, M. Cadossi, N. Baldini and S. Giannini, Clin. Orthop., 2008, 466(3), 700–707.

[28] S. Smith, D. Dowson and A. Goldsmith, Proc. Inst. Mech. Eng. H J. Eng. Med., 2001, 215(J5), 483–493.

[29] M. Clarke, P. Lee, A. Arora and R. Villar, J. Bone Joint Surg. Br, 2003, 85(6), 913–917.

[30] M. Bernstein, A. Walsh, A. Petit, D. Zukor and J. Antoniou, Clin. Or- thop., 2011, 469(6), 1642–1650.

[31] W. Witzleb, J. Ziegler, F. Krummenauer, V. Neumeister and K. Guen- ther, Acta Orthop., 2006, 77(5), 697–705.

[32] J. Antoniou, D. Zukor, F. Mwale, W. Minarik, A. Petit and O. Huk, J. Bone Joint Surg. Am., 2008, 90(Supplement 3), 142–148.

[33] J. Daniel, H. Ziaee, A. Salama, C. Pradhan and D. J. W. McMinn, J. Bone Joint Surg. Br, 2006, 88(4), 443–448.

[34] A. Moroni, L. Savarino, M. Hoque, M. Cadossi and N. Baldini, Clin. Orthop., 2011, 469(1), 180–187.

[35] R. De Haan, C. Pattyn, H. Gill, D. Murray, P. Campbell and K. De Smet, J. Bone Joint Surg. Br, 2008, 90(10), 1291–1297.

[36] S. Suñer, J. Tipper and N. Emami, Tribol. - Mater. Surf. Interfaces, 2012, 6(2), 39–52.

[37] I. Papageorgiou, C. Brown, R. Schins, S. Singh, R. Newson, S. Davis, J. Fisher, E. Ingham and C. Case, Biomaterials, 2007, 28(19), 2946– 2958.

[38] Y. Kwon, Z. Xia, S. Glyn-Jones, D. Beard, H. Gill and D. Murray, Biomed. Mater., 2009, 4(2), 025018.

[39] M. Germain, A. Hatton, S. Williams, J. Matthews, M. Stone, J. Fisher and E. Ingham, Biomaterials, 2003, 24(3), 469–479.

[40] H. Willert, G. Buchhorn, A. Fayyazi, R. Flury, M. Windler, G. Köster and C. Lohmann, J. Bone Joint Surg. Am., 2005, 87(1), 28–36. REFERENCES 73

[41] N. Hallab, K. Merritt and J. Jacobs, J. Bone Joint Surg. Am., 2001, 83(3), 428–428. [42] N. Hallab, M. Caicedo, R. Epstein, K. McAllister and J. Jacobs, J. Biomed. Mater. Res., 2010, 92(2), 667–682. [43] V. Ng, A. Lombardi, K. Berend, M. Skeels and J. Adams, Clin. Orthop., 2011, 469(2), 523–529. [44] P. Campbell, A. Shimmin, L. Walter and M. Solomon, J. Arthroplasty, 2008, 23(7), 1080–1085. [45] D. Neumann, C. Thaler, W. Hitzl, M. Huber, T. Hofstädter and U. Dorn, J. Arthroplasty, 2010, 25(5), 700–708. [46] A. Carrothers, R. Gilbert, A. Jaiswal and J. Richardson, J. Bone Joint Surg. Br, 2010, 92(10), 1344–1350. [47] T. Fujishiro, D. Moojen, N. Kobayashi, W. Dhert and T. Bauer, Clin. Orthop., 2011, 469(4), 1127–1133. [48] C. Engh, H. Ho and C. Engh, Clin. Orthop., 2010, 468(2), 406–412. [49] P. Korovessis, G. Petsinis, M. Repanti and T. Repantis, J. Bone Joint Surg. Am., 2006, 88(6), 1183–1191. [50] J. Thyssen, S. Jakobsen, K. Engkilde, J. Johansen, K. Søballe and T. Menné, Acta Orthop., 2009, 80(6), 646–652. [51] P. Thomas, L. Braathen, M. Dörig, J. Auböck, F. Nestle, T. Werfel and H. Willert, Allergy, 2009, 64(8), 1157–1165. [52] W. Witzleb, U. Hanisch, N. Kolar, F. Krummenauer and K. Guenther, Acta Orthop., 2007, 78(2), 211–220. [53] H. Pandit, S. Glyn-Jones, P. McLardy-Smith, R. Gundle, D. Whitwell, C. Gibbons, S. Ostlere, N. Athanasou, H. Gill and D. Murray, J. Bone Joint Surg. Br, 2008, 90(7), 847–851. [54] S. Glyn-Jones, H. Pandit, Y. Kwon, H. Doll, H. Gill and D. Murray, J. Bone Joint Surg. Br, 2009, 91(12), 1566–1574. [55] G. Grammatopolous, H. Pandit, Y.-M. Kwon, R. Gundle, P. McLardy- Smith, D. Beard, D. Murray and H. Gill, J. Bone Joint Surg. Br, 2009, 91(8), 1019–1024. 74 REFERENCES

[56] A. Malviya and J. Holland, Acta Orthop. Belg., 2009, 75(4), 477–483.

[57] G. Mahendra, H. Pandit, K. Kliskey, D. Murray, H. Gill and N. Athana- sou, Acta Orthop., 2009, 80(6), 653–659.

[58] H. Pandit, M. Vlychou, D. Whitwell, D. Crook, R. Luqmani, S. Ostlere, D. Murray and N. Athanasou, Virchows Arch., 2008, 453(5), 529–534.

[59] H. Sieber, C. Rieker and P. Köttig, J. Bone Joint Surg. Br, 1999, 81(1), 46–50.

[60] J. Daniel, P. Pynsent and D. McMinn, J. Bone Joint Surg. Br, 2004, 86(2), 177–184.

[61] S. Saito, J. Ryu, M. Watanabe, T. Ishii and K. Saigo, J. Arthroplasty, 2006, 21(8), 1105–1110.

[62] Y. Park, Y. Moon, S. Lim, J. Yang, G. Ahn and Y. Choi, J. Bone Joint Surg. Am., 2005, 87(7), 1515–1521.

[63] F. Antony and C. Holden, Contact Dermatitis, 2003, 48(1), 49–50.

[64] J. Zustin, M. Hahn, M. Morlock, W. Rüther, M. Amling and G. Sauter, Skeletal Radiol., 2010, 39(8), 747–756.

[65] V. Milavec-Puretic,´ D. Orlic´ and A. Marušic,´ Arch. Orthop. Trauma Surg., 1998, 117, 383–386.

[66] R. Shrivastava, R. Upreti, P. Seth and U. Chaturvedi, FEMS Immunol. Med. Microbiol., 2002, 34(1), 1–7.

[67] D. Beyersmann and A. Hartwig, Toxicol. Appl. Pharmacol., 1992, 115(1), 137–145.

[68] I. Papageorgiou, Z. Yin, D. Ladon, D. Baird, A. Lewis, A.Sood, R. New- son, I. Learmonth and C. Case, Mutat. Res., 2007, 619(1-2), 45 – 58.

[69] A. Davies, A. Sood, A. Lewis, R. Newson, I. Learmonth and C. Case, J. Bone Joint Surg. Br, 2005, 87(10), 1439–1444.

[70] D. Ladon, A. Doherty, R. Newson, J. Turner, M. Bhamra and C. Case, J. Arthroplasty, 2004, 19(8), 78–83.

[71] A. Doherty, R. Howell, L. Ellis, I. Bisbinas, I. Learmonth, R. Newson and C. Case, J. Bone Joint Surg. Br, 2001, 83(7), 1075–1081. REFERENCES 75

[72] R. Tharani, F. Dorey and T. Schmalzried, J. Bone Joint Surg. Am., 2001, 83(5), 774–780.

[73] T. Visuri, P. Pulkkinen, P. Paavolainen and E. Pukkala, Acta Orthop., 2010, 81(1), 77–81.

[74] T. Visuri, E. Pukkala, P. Paavolainen, P. Pulkkinen and E. Riska, Clin. Orthop., 1996, 329, S280–289.

[75] W. Gillespie, D. Henry, D. O’Connell, S. Kendrick, K. McInneny, L. Derby et al., Clin. Orthop., 1996, 329, S290–S296.

[76] O. Nyrén, J. McLaughlin, G. Gridley, A. Ekbom, O. Johnell, J. Frau- meni and H. Adami, J. Natl. Cancer Inst., 1995, 87(1), 28–33.

[77] P. Wagner, H. Olsson, L. Lidgren, O. Robertsson and J. Ranstam, Eur. J. Cancer, 2011, 47(7), 1061–1071.

[78] T. Murakami, H. Higaki, Y. Sawae, N. Ohtsuki, S. Moriyama and Y. Nakanishi, Proc. Inst. Mech. Eng. H J. Eng. Med., 1998, 212(H1), 23–35, 1997 World Tribology Congress, London, England, SEP 08-12, 1997.

[79] J. Delecrin, M. Oka, S. Takahashi, T. Yamamuro and T. Nakamura, Clin. Orthop., 1994, (307), 240–249.

[80] F. Liu, Z. Jin, C. Rieker, F. Hirt, P. Roberts and P. Grigoris, J. Bone Joint Surg. Br, 2006, 88-B(SUPPIII), 387–387.

[81] D. Dowson, C. Hardaker, M. Flett and G. Isaac, J. Arthroplasty, 2004, 19(8), 118–123.

[82] D. Dowson, C. Hardaker, M. Flett and G. Isaac, J. Arthroplasty, 2004, 19(8), 124–130.

[83] J. Yang, B. Shen, Z. Zhou, F. Pei and P. Kang, J. Arthroplasty, 2011, 26(1), 65–70, 70.e1.

[84] J. Daniel, H. Ziaee, C. Pradhan, P. Pynsent and D. McMinn, J. Bone Joint Surg. Br, 2007, 89(2), 169–173.

[85] J. Daniel, H. Ziaee, C. Pradhan and D. McMinn, J. Bone Joint Surg. Br, 2009, 91(2), 176–179. 76 REFERENCES

[86] M. Wimmer, J. Loos, R. Nassutt, M. Heitkemper and A. Fischer, Wear, 2001, 250(1-12), 129 – 139, 13th International Conference on Wear of Materials.

[87] M. Wimmer, C. Sprecher, R. Hauert, G. Täger and A. Fischer, Wear, 2003, 255(7), 1007–1014.

[88] T. Hanawa, Mater. Sci. Eng., C, 2004, 24(6-8), 745 – 752.

[89] A. Hodgson, S. Kurz, S. Virtanen, V. Fervel, C.-O. Olsson and S. Mis- chler, Electrochim. Acta, 2004, 49(13), 2167 – 2178.

[90] I. Gotman, J. Endourol., 1997, 11(6), 383–389, 1st International Sym- posium on Urological Stents, Jerusalem, Israel, OCT, 1996.

[91] J. Jacobs, J. Gilbert and R. Urban, J. Bone Joint Surg. Am., 1998, 80A(2), 268–282.

[92] G. C. F. Clark and D. F. Williams, J. Biomed. Mater. Res., 1982, 16(2), 125–134.

[93] A. Muñoz and S. Mischler, J. Electrochem. Soc., 2007, 154(10), C562– C570.

[94] A. Sargeant and T. Goswami, Mater. Design, 2007, 28(1), 155–171.

[95] D. R. Jackson, S. Omanovic and S. G. Roscoe, Langmuir, 2000, 16(12), 5449–5457.

[96] X. Cheng and S. G. Roscoe, Biomaterials, 2005, 26(35), 7350–7356.

[97] M. A. Khan, R. L. Williams and D. F. Williams, Biomaterials, 1999, 20(7), 631–637.

[98] A. C. Lewis, M. R. Kilburn, P. J. Heard, T. B. Scott, K. R. Hallam, G. C. Allen and I. D. Learmonth, J. Orthop. Res., 2006, 24(8), 1587–1596.

[99] Y. Yan, A. Neville, D. Dowson and S. Williams, Tribol. Int., 2006, 39(12), 1509–1517.

[100] Y. Yan, A. Neville, D. Dowson, S. Williams and J. Fisher, Proc. Inst. Mech. Eng. J J. Eng. Tribol., 2008, 222(3), 483–492.

[101] D. Sun, J. A. Wharton, R. J. K. Wood, L. Ma and W. M. Rainforth, Tribol. Int., 2009, 42(1), 99–110. REFERENCES 77

[102] M. Stack, J. Rodling, M. Mathew, H. Jawan, W. Huang, G. Park and C. Hodge, Wear, 2010, 269(5), 376–382.

[103] Y. Yan, A. Neville, D. Dowson, S. Williams and J. Fisher, Wear, 2009, 267(5), 683–688.

[104] Y. Yan, A. Neville, D. Dowson, S. Williams and J. Fisher, Proc. Inst. Mech. Eng. J J. Eng. Tribol., 2009, 223(3), 303–309.

[105] Y. Yan, A. Neville, D. Dowson, S. Williams and J. Fisher, Proc. Inst. Mech. Eng. H J. Eng. Med., 2010, 224(11), 1267–1273.

[106] Y. Yan, A. Neville, D. Dowson, S. Williams and J. Fisher, Proc. Inst. Mech. Eng. J J. Eng. Tribol., 2010, 224(9), 997–1006.

[107] P. Ponthiaux, F. Wenger, D. Drees and J. Celis, Wear, 2004, 256(5), 459–468.

[108] Y. Yan, A. Neville and D. Dowson, Wear, 2007, 263(7), 1105–1111.

[109] N. Hallab, A. Skipor and J. Jacobs, J. Biomed. Mater. Res., Part A, 2003, 65(3), 311–318.

[110] M. Wimmer, A. Fischer, R. Büscher, R. Pourzal, C. Sprecher, R. Hauert and J. Jacobs, J. Orthop. Res., 2010, 28(4), 436–443.

[111] Y. Liao, R. Pourzal, M. A. Wimmer, J. J. Jacobs, A. Fischer and L. D. Marks, Science, 2011, 334(6063), 1687–1690.

[112] L. Gao, F. Wang, P. Yang and Z. Jin, Med. Eng. Phys., 2009, 31(6), 720–729.

[113] W.-Z. Wang, Z. M. Jin, D. Dowson and Y. Z. Hu, Proc. Inst. Mech. Eng. J J. Eng. Tribol., 2008, 222(3), 493–501.

[114] W.-Z. Wang, F.-C. Wang, Z.-M. Jin, D. Dowson and Y.-Z. Hu, Proc. Inst. Mech. Eng. J J. Eng. Tribol., 2009, 223(7), 1073–1082.

[115] Z. Jin and D. Dowson, Proc. Inst. Mech. Eng. C J. Mech. Eng. Sci., 1999, 213(4), 355–370.

[116] D. Jalali-Vahid and Z. Jin, Proc. Inst. Mech. Eng. C J. Mech. Eng. Sci., 2002, 216(4), 409–420.

[117] D. Meyer and T. J.A., J. Biomech. Eng., 2003, 125, 777–784. 78 REFERENCES

[118] F. Wang, C. Brockett, S. Williams, I. Udoﬁa, J. Fisher and Z. Jin, Phys. Med. Biol., 2008, 53(5), 1277–1293.

[119] D. Dowson, Proc. Inst. Mech. Eng. H J. Eng. Med., 2006, 220(2), 161– 171.

[120] S. Mischler, S. Debaud and D. Landolt, J. Electrochem. Soc., 1998, 145(3), 750–758.

[121] J. Jiang, M. Stack and A. Neville, Tribol. Int., 2002, 35(10), 669–679. Paper B

Corrosion and Tribocorrosion of Hafnium in Simulated Body Fluids

B.1. ABSTRACT 81

Submitted for journal publication

Corrosion and Tribocorrosion of Hafnium in Simulated Body Fluids

J. Rituerto Sin1, A. Neville2 and N. Emami1 1Luleå University of Technology, Division of Machine Elements, Luleå, SE-971 87 Sweden 2Institute of Engineering Thermoﬂuids, Surfaces and Interfaces, School of Mechanical Engineering, University of Leeds, Leeds LS2 9JT UK

B.1 Abstract

Hafnium is a passive metal with good biocompatibility and osteogenesis, however, little is known about its resistance to wear and corrosion in biological environments. The corrosion and tribocorrosion behaviour of hafnium and commercially pure (CP) titanium in simulated body ﬂuids were investigated using electrochemical techniques. Cyclic polarisation scans and open circuit potential measurements were performed in 0.9% NaCl solution and 25% bovine calf serum solution in order to assess the effect of organic species on the corrosion behaviour of the metal. A pin-on-plate conﬁguration tribometer and a three electrode electrochemical cell were integrated to investigate the tribocorrosion performance of the studied materials. The results showed that hafnium has good corrosion resistance. The corrosion density currents measured in its passive state were lower than those measured in the case of CP titanium, however, it showed a higher tendency to suffer from localised corrosion, which was more acute when imperfections were present on the surface. The electrochemical breakdown of the oxide layer was retarded in the presence of proteins. Tribocorrosion tests showed that hafnium has the ability to quickly repassivate after the oxide layer was damaged; however, it showed higher volumetric loss than CP titanium in equivalent wear-corrosion conditions.

Keywords: Hafnium, Titanium, Biomaterial, Corrosion, Tribocorrosion 82 PAPER B. CORROSION AND TRIBOCORROSION OF HAFNIUM

B.2 Introduction

Hafnium is a metal included in the group 4 in the periodic table, together with titanium and zirconium. The metal was first identified by Dirck Coster and Georges de Hevesy in Copenhagen in 1923 [1] and owes its name to ’Hafnia’, the Latin name for Copenhagen. Hafnium is always found associated with zirconium in ores and is primarily found in zircon, with a ratio Hf/Zr of about 2.5% [2]. Due to a number of interesting properties such as high ductility and strength, as well as resistance to corrosion and mechanical damage [3], it has attracted interest for various applications. For instance, it is used as a control material for nuclear reactors and as an alloying element in some superalloys used in aircrafts engines [4]. In 1984, Marcel Pourbaix included hafnium among the group of metals to be considered for surgical implants due to the passive state presented by the metal. However, due to the lack of information about its toxicity in the human body at that time it was discarded from the final list of 13 metals to be theoretically considered [5]. The passive layer found on hafnium’s surface, mainly formed of HfO2, is highly stable in neutral solutions [6]. More recently, the properties of hafnium as an implant material have been investigated [7, 8]. Matsuno et al. [7] implanted a number of refractory materials (titanium, hafnium, niobium, tantalum and rhenium) in subcutaneous tissue and femoral bone marrow of rats, and performed histological observation and elemental mapping in order to evaluate the biocompatibility and osteogenesis of such materials. The results of the in vivo study showed that hafnium had good biocompatibility and osteogenesis. A different study on the tissue response to hafnium [8] showed similar response in soft and hard tissues in two different animal species (rat and rabbit), which suggests that hafnium could be an interesting material for biomedical applications. Hafnium has also been investigated as an alloying element for titanium alloys. Titanium alloys are commonly used for biomedical applications such as dental implants, bone plates and hip and knee implants. Titanium-6%Aluminum- 4%Vanadium (Ti-6Al-4V) is usually the material of choice. However concerns regarding the use of aluminium -which has been pointed out as a growth inhibitor of bone and possible cause of Alzheimer’s disease [9]- and vanadium -which can lead to cardiac and renal dysfunction [10]- enhance the need to investigate new alloying elements. Different Ti-Hf binary alloys have been reported in the literature. Wrought Ti-Hf alloys have shown a gradually increase in yield and tensile strength with increased amount of hafnium, up to 40%, at the same time the ductility decreased [11]. More recent studies also showed B.3. MATERIALS AND METHODS 83 an increased yield and tensile strength as well as bulk hardness for cast Ti-Hf alloys [12]. The electrochemical behaviour of this type of alloys is similar to pure titanium, showing excellent corrosion resistance [13]. Titanium-Niobium-Hafnium-Zirconium (Ti-Nb-Hf-Zr) alloys have also been investigated [14]. These alloys have shown a low elastic modulus which is beneficial in order to reduce the stress shielding effect and to enhance bone growth. It has been also shown that cold work can be used to decrease the elastic modulus of this type of alloys, reaching values close to the elastic modulus of cortical bone [15]. To date, the behaviour of pure hafnium in a biological environment has not been studied in great depth. Little is known about the resistance of the passive layer under wear-corrosion conditions and the effect of proteins on its corrosion and tribocorrosion behaviour. Thus, there exists a need for further studies of hafnium under biological conditions in order to determine the suitability of this material for biomedical applications. The present work studies the corrosion resistance of hafnium in simulated body fluids as well as the effect of wear on the system.

B.3 Materials and Methods

B.3.1 Preparation of specimens and electrolyte solutions Commercially pure titanium (Grade 2) and commercially pure hafnium rods (GfE Metalle und Materialien GmbH, Germany) were used in this study. Plates 6 mm thickness by 25 mm diameter were prepared from the initial rods. Spec- imens were cast in an epoxy mount and the epoxy-metal borders were sealed with silicon to avoid the presence of crevices. The plates were polished with silicon carbide (SiC) paper up to 1200 grit followed by a 9 µm and a 3 µm diamond paste. Finish polishing was accomplished with a colloidal silica suspension. Specimens were cleaned and rinsed with distilled water and ethanol before they were assembled into the measurement cell. Two different surface finishes were investigated in the case of hafnium: a first group of specimens were studied immediately after polishing; a second group of specimens were scratched after polishing, to assess the effect of surface imperfections on the corrosion resistance of hafnium. In order to simulate body fluids, 25% and 50% bovine calf serum (Harlam Sera-LabÂo,˝ UK) solutions were employed. In addition, 0.9% NaCl solution was used to isolate the effect of the organic species (proteins, amino acids, etc.) from the saline environment. All tests were conducted at 37 ± 1◦C. 84 PAPER B. CORROSION AND TRIBOCORROSION OF HAFNIUM

B.3.2 Electrochemical tests

All electrochemical measurements were carried out using an Autolab PG- STAT302N potentiostat (Metrohm Autolab, The Netherlands). A three electrode electrochemical set-up was used. A platinum wire was used as a counter electrode, a Ag/AgCl electrode (3 M KCl) was used as a reference electrode (196 mV vs SHE) and the studied material acted as a working electrode. All potentials are given with respect to the Ag/AgCl electrode. Specimens were immersed in the solution at open circuit potential for 60 min to stabilize them. Cyclic polarization tests were performed at 2 mV s-1 scan rate. The current density was limited to a maximum of 500 µA cm-2. All tests were repeated at least three times to ensure repeatability. Optical microscopy (Olympus Vanox- T, Olympus Optical, Japan) was used to analyse the surface of the specimens before and after measurements.

B.3.3 Tribocorrosion tests

A tribocorrosion cell compromising of a three electrode electrochemical cell and a reciprocating tribological tester was used. All electrochemical measurements were carried out using a potentiostat VersaSTAT 4 (Princeton Applied Research, USA). The three-electrode cell configuration is the same as the earlier described, with an Ag/AgCl electrode as the reference electrode, a Pt wire as counter electrode and the specimen as the working electrode. The sliding wear tests were performed in a reciprocating wear tester that conforms to ASTM G133. The studied materials were rubbed against a silicon nitride ball with 12 mm diameter. A normal load of 40 N was used. The frequency was controlled at 1 Hz and the stroke length was set at 10 mm. Specimens were passivated in 30% HNO3 for 30 minutes before each test. The mass of the specimens was gravimetrically determined by using a digital microbalance (Mettler Toledo AT201, UK) with 0.01 mg resolution. Each pin was weighed before and after each test at least five times to ensure repeatability. Specimens were immersed in the selected solution in static conditions before the onset of the tribological test and the OCP was monitored for 60 minutes. After that the tribological test was started and the OCP was measured for 60 minuntes. Finally, the sliding was stopped and the OCP was measured for 60 minutes to study the repassivation kinetics of the materials in different environments. 3D optical profilometry (Wyko 1100NT, Veeco Instruments, USA) was used to study the cross-section profile of the wear tracks after the tribocorrosion tests. B.4. RESULTS 85

Table B.1: Breakdown potential of hafnium in 0.9% NaCl and 25% serum. 0.9% NaCl 25% serum Ebreakdown (V) 1.38 ± 0.11 2.33 ± 0.18 ±: Standard deviation

3.0 3.0 Hf - 0.9% NaCl Hf - 0.9% NaCl 2.5 2.5 Hf - 25% serum Hf - 25% serum 2.0 2.0

1.5 1.5

1.0 1.0

0.5 0.5

0.0 0.0 0.0E+00 5.0E-04 1.0E-03 1.5E-03 2.0E-03 2.5E-03 3.0E-03 0.00E+00 1.25E-04 2.50E-04 3.75E-04 5.00E-04 6.25E-04 -0.5 -0.5

Potential (V) vs Ag/AgCl Potential (V) vs -1.0 Ag/AgCl Potential (V) vs -1.0 Current Density (A cm-2) Current Density (A cm-2) (a) (b)

Figure B.1: Cyclic polarisation curve of (a) mirror polished hafnium and (b) scratched hafnium.

B.4 Results

B.4.1 Electrochemical behaviour

The potentiodynamic polarization curves of hafnium in 0.9% NaCl solution and 25% serum solution are shown in Figure B.1(a). A sudden increase of the current density was observed in both solutions due to the breakdown of the passive layer. The breakdown potential increased in presence of proteins, as shown in Table B.1. Once the current density reached the current limit (500 µA cm-2), the potential slope was reversed, however the current density kept increasing after the reversal point was reached, showing typical pit propagation behaviour. The pit propagation curve was larger in 0.9% NaCl solution. The repassivation potential in the case of 25% serum solution was slightly higher than in the saline solution. As shown in Figure B.1(b), the breakdown potential showed a large decrease when the surface of the material was scratched. Formation of pits due to localised corrosion was observed under optical microscopy. The pits observed on the polished specimens (Figure B.2(a) and B.2(b)) were larger than in the case of the scratched surfaces, where, small pits were preferentially formed on the surroundings of the surface imperfections (Figure B.2(c) and B.2(d)). 86 PAPER B. CORROSION AND TRIBOCORROSION OF HAFNIUM

(a) (b)

Figure B.2: Corrosion pits observed on the surface of hafnium after cyclic polarisation tests on the mirror polished surface in (a) 0.9% NaCl and (b) 25% serum and on the scratched surface in (c) 0.9% NaCl and (d) 25% serum. B.4. RESULTS 87

! " ! # ! # "

Figure B.3: Anodic polarisation curves of hafnium and CP titanium in 0.9% NaCl and 25% serum.

Table B.2: Anodic current density of hafnium and CP titanium in the passive region at different applied potentials. Hafnium CP titanium 0.9% NaCl 25% serum 0.9% NaCl 25% serum Potential (mV) Passive Current Density (µAcm−2) 0 4.26 ± 0.18 5.65 ± 0.02 5.95 ± 1.52 11.99 ± 0.24 250 5.36 ± 0.03 6.08 ± 0.05 9.66 ± 0.79 11.17 ± 0.16 500 6.25 ± 0.25 6.43 ± 0.01 10.21 ± 0.02 10.90 ± 0.31 750 6.97 ± 0.38 6.78 ± 0.19 10.31 ± 0.49 10.46 ± 0.50 1000 7.44 ± 0.38 7.15 ± 0.10 10.2 ± 0.96 10.28 ± 0.07 1250 7.82 ± 0.07 7.34 ± 0.06 10.8 ± 0.31 10.33 ± 0.28 ±: Standard deviation

The anodic polarisation curve was extracted from the cyclic polarisation tests. Figure B.3 shows the anodic polarisation curves of hafnium and CP titanium in 0.9% NaCl and 25% serum solution, which could be divided in four zones: at potentials below the corrosion potential (Ecorr), the current density was dominated by the cathodic reaction; at potentials around the Ecorr there was a transition from cathodic to anodic currents; when the potential was increased above the Ecorr, both materials showed a passive plateau where a stable oxide layer was formed on their surface; ﬁnally, when the potential was further increased the passive ﬁlm was disrupted and the current rapidly increased. In the case of CP titanium, breakdown of the oxide layer was not observed, the passive layer remained stable at high potentials. The current densities of hafnium showed 88 PAPER B. CORROSION AND TRIBOCORROSION OF HAFNIUM

Table B.3: Tafel extrapolation of the corrosion potential and corrosion current density in 0.9% NaCl and 25% serum. Hafnium CP titanium 0.9% NaCl 25% serum 0.9% NaCl 25% serum -0.638 -0.881 -0.415 -0.613 E (V) corr ± 0.016 ± 0.045 ± 0.011 ± 0.051 163.9 15.3 395.9 146.0 i (nA cm−2) corr ± 40.9 ± 3.5 ± 114.9 ± 32.0 ±: Standard deviation

Hafnium Titanium 0.0 0.0 0 3600 7200 10800 0 3600 7200 10800 -0.2 -0.2

-0.4 -0.4

-0.6 -0.6

-0.8 -0.8

-1.0 -1.0

-1.2 0.9% NaCl -1.2 0.9% NaCl

OCP (V) vs Ag/AgCl (V) vs OCP 25% serum Ag/AgCl (V) vs OCP 25% serum -1.4 -1.4 50% serum 50% serum -1.6 -1.6 Time (s) Time (s) (a) (b)

Figure B.4: Evolution of the OCP during tribocorrosion test on (a) hafnium and (b) CP titanium. slightly lower current density values than titanium in its passive state (Table B.2) before the breakdown potential was reached. The values of corrosion potential (Ecorr) and corrosion current density (icorr) of titanium and hafnium in 0.9% NaCl solution and 25% serum solution extracted from the Tafel plots ﬁt- ted to the potentiodynamic curves are shown in Table B.3. Hafnium showed a lower Ecorr than titanium in both solutions. The icorr observed for hafnium was slightly lower than that for CP titanium in both 0.9% NaCl solution and 25% serum. Both materials showed lower icorr in 25% serum.

B.4.2 Tribocorrosion behaviour The evolution of the OCP value during the tribocorrosion tests is shown in Fig- ure B.4. Stable passive OCP values were observed for the ﬁrst 60 min, when the specimens were immersed in the electrolyte solution. With the onset of the reciprocating test a shift of the OCP value towards more active values was B.4. RESULTS 89

(a) (b) (c)

Figure B.5: Evolution of the OCP of hafnium and CP titanium after mechanical damaging ceased in (a) 0.9% NaCl, (b) 25% serum and (c) 50% serum solution. observed, indicating that the passive layer was being mechanically damaged. When the mechanical damaging ceased the OCP rapidly increased, indicating that the oxide layer was rebuilt on the surface. The evolution of the OCP of hafnium and CP titanium after mechanical depassivation ceased is shown in Figure B.5. The percentage of repassivataion was calculated according to Equation (1). E − E Repassivation(%)= t wear (B.1) Epassive − Ewear

Where Et is the OCP value at a time t; Ewear is the OCP value during wear- corrosion before the mechanical damage ceased; and Epassive is the OCP value 90 PAPER B. CORROSION AND TRIBOCORROSION OF HAFNIUM

0.007 Hafnium 0.006

) Titanium 3 0.005 (cm 0.004 Loss

0.003

0.002 Volumetric

0.001

0 0.9%NaCl 25%serum 50%serum

Figure B.6: Total volumetric loss of hafnium and CP titanium after the tribocorrosion test in 0.9% NaCl, 25% serum and 50% serum.

0.1 0.1

0.05 0.05

0 0 00.511.522.533.544.55 00.511.522.533.544.55 -0.05 -0.05

-0.1 -0.1 mm mm -0.15 -0.15

-0.2 -0.2 Hf - 0,9% NaCl CP Ti - 0,9% NaCl -0.25 Hf - 25% serum -0.25 CP Ti - 25% serum Hf - 50% serum CP Ti - 50% serum -0.3 mm -0.3 mm (a) (b)

Figure B.7: Wear scar cross-section profiles measured after the tribocorrosion tests on (a) hafnium and (b) CP titanium. before the mechanical damaging starts. After the first quick increase, the potential kept rising at a lower rate due to the growth of the protective film. The OCP increased slightly faster in the case of hafnium in all fluids. The measured volumetric losses of hafnium and titanium in 0.9% NaCl solution, 25% serum and 50% serum are shown in Figure B.6. Hafnium showed higher volumetric loss in 0.9% NaCl solution and 25% serum solution. The volumetric loss was generally higher for 25% serum than for 0.9% NaCl solution, however, it did not increase in the case of 50% serum in comparison with 25% serum. The depth of the wear scars measured for hafnium was larger in the case of hafnium (Figure B.7(a)) than for CP titanium (Figure B.7(b)) in all solutions. B.5. DISCUSSION 91

B.5 Discussion

B.5.1 Corrosion resistance of hafnium The pH of the human body is normally around 7.4, however, it can be affected by different factors such as infections or surgery, where the pH can vary from values around 5.6 up to 9 [5]. A preliminary study of the potential-pH curves of hafnium and titanium showed that these two materials are in passive state in the body’s pH range [5]. Hafnium is spontaneously passivated in aqueous environments by forming a hafnium oxide layer on the surface [6]. Titanium is well known to be in the passive state in biological environments. A titanium oxide layer provides this material with excellent corrosion resistance. Titanium and hafnium are generally passivated by the reaction with water according to the equation (1) and (2) respectively. + − Ti+ 2H2O ↔ TiO2 + 4H + 4e (B.2) + − Hf+ 2H2O ↔ HfO2 + 4H + 4e (B.3) In simulated body ﬂuids, such as the studied solutions, both titanium and hafnium formed a stable passive layer and OCP values were in the passive domain. The oxide layer formed on hafnium has been shown to be very protective, showing slightly lower current densities than those for titanium. Both materials also showed low corrosion current density in the presence of organic species. The icorr values extracted from the Tafel plots showed lower corrosion currents densities in 25% serum solution, indicating that proteins might be acting as a cathodic inhibitor [16], however, no signiﬁcant difference was observed in the corrosion current densities measured in the passive region in the saline and in the serum solutions.

Cyclic polarisation tests revealed that hafnium tends to suffer from localised corrosion at higher potentials. The results showed that the presence of organic species increased the breakdown potential of the ﬁlm, which could be due to the positive charge of proteins at the pH of the test solutions (7.4) [17]. In the polarisation scan, specimens were anodically polarised, therefore the negatively charged species tended to be attracted towards the surface of the specimen. Proteins would compete with chlorides to adsorb on the surface, which would retard the formation of pits. Similar effects have been reported for AISI 316L stainless steel [18]. The breakdown potential measured in both 0.9% serum and 25% serum solution exceeded the anodic oxygen evolution potential, however this reaction could not be detected, due to the insulating 92 PAPER B. CORROSION AND TRIBOCORROSION OF HAFNIUM

effect of HfO2 for anodic reactions [19]. The pit propagation curve was also limited in the presence of organic species. Proteins would be attracted towards the region were positively charged ions were released [18]. The presence of proteins could be act as a diffusion barrier; therefore the current density values reached when the pits were formed in serum solution were lower than in the case of saline solution. A strong correlation between the surface ﬁnish and resistance to localised corrosion was also observed. When surface imperfections were present on the surface of hafnium, its resistance to localised corrosion was greatly decreased, and corrosion pits were formed around the surface imperfections.

B.5.2 Tribocorrosion of hafnium As discussed earlier, both hafnium and titanium show a passive state when immersed in aqueous solution. The passive state was also observed in the presence of organic species, which provided them with excellent corrosion resistance. The mechanical damage to which the materials were exposed caused damage to the passive layer and the bulk material to be in contact with the media, therefore the OCP dropped towards more active values. When the mechanical damage ceased, both materials had the ability to repassivate, and the oxide layer was quickly reformed on their surface, which could be observed by a quick increase on the OCP. Hafnium showed a slightly faster repassivation that was observed by a faster increase of the OCP in all solutions. The presence of organic species slowed down the reformation of the passive layer in both cases. The volumetric loss measured in the case of hafnium was higher in 0.9% NaCl and 25% serum under the studied conditions. This was inﬂuenced by the higher elastic modulus of hafnium (138MPa [20]) as compared to CP titanium (102.7 MPa [21]), which yields an initial Hertzian contact pressure of approximately 850 MPa in the case of hafnium and approximately 750 MPa in the case of CP titanium in this study. However, the volumetric loss decreased in 50% serum for both materials in comparison with 25% serum solution, which suggested that the increase of the protein content of this solution may improve the lubrication of the system [22, 23].

B.6 Conclusions

Hafnium has been pointed out as an interesting material for biomedical applications due to its good biocompatibility and osteogenesis [8]. However its B.7. ACKNOWLEDGEMENTS 93 corrosion and tribocorrosion behaviour in the biological environment have not been fully studied. In the present study, the corrosion and tribocorrosion behaviour of hafnium have been investigated using CP titanium as the reference material.

• Hafnium showed a passive state in the presence of proteins. The oxide layer formed provided high protection to corrosion.

• Hafnium had a higher tendency to suffer from localised corrosion than CP titanium. The presence of proteins retarded the formation of pits and reduced their propagation.

• Mechanical damage could disrupt the passive layer of hafnium prevent- ing its protective effect. However, the oxide layer was quickly reformed on the surface when the damaging ceased. Hafnium showed a slightly faster repassivation than titanium.

• The wear rate shown by hafnium was higher in comparison to CP titanium in 0.9% NaCl solution and 25% solution under the studied conditions.

B.7 Acknowledgements

J. Rituerto would like to acknowledge the Swedish Research of Tribology for supporting the visit to the University of Leeds, UK. This work was funded by the Centrum för Medicinsk Teknik och Physic (CMTF) and Kempe Stif- telserna.

References

[1] D. Coster and G. de Hevesy, Nature, 1923, 111, 79.

[2] R. Tricot, J. Nucl. Mater., 1992, 189(3), 277–288.

[3] E. Cerreta, C. Yablinsky, G. Gray, S. Vogel and D. Brown, Mater. Sci. Eng., A, 2007, 456(1), 243–251.

[4] O. Levy, G. L. Hart and S. Curtarolo, Acta Mater., 2010, 58(8), 2887– 2897.

[5] M. Pourbaix, Biomaterials, 1984, 5(3), 122–134. 94 REFERENCES

[6] W. A. Badawy and F. M. Al-Kharaﬁ, J. Mater. Sci., 1999, 34, 2483–2491.

[7] H. Matsuno, A. Yokoyama, F. Watari, M. Uo and T. Kawasaki, Biomate- rials, 2001, 22(11), 1253–1262.

[8] S. Mohammadi, M. Esposito, M. Cucu, L. E. Ericson and P. Thomsen, J. Mater. Sci.: Mater. Med., 2001, 12(7), 603–611.

[9] L. Tomljenovic, J. Alzheimers Dis., 2011, 23(4), 567–598.

[10] N. J. Hallab and J. J. Jacobs, Bull. NYU Hosp. Jt. Dis., 2009, 67(2), 182– 188.

[11] A. Imgram, D. Williams and H. Ogden, J. Less-Common Met., 1962, 4(3), 217–225.

[12] H. Sato, M. Kikuchi, M. Komatsu, O. Okuno and T. Okabe, J. Biomed. Mater. Res., Part B, 2005, 72(2), 362–367.

[13] Z. Cai, M. Koike, H. Sato, M. Brezner, Q. Guo, M. Komatsu, O. Okuno and T. Okabe, Acta Biomater., 2005, 1(3), 353–356.

[14] M. González, J. Peña, J. Manero, M. Arciniegas and F. Gil, J. Mater. Eng. Perform., 2009, 18(5), 490–495.

[15] M. González, F. Gil, J. Manero and J. Peña, J. Mater. Eng. Perform., 2011, 20(4), 653–657.

[16] A. Muñoz and S. Mischler, J. Electrochem. Soc., 2007, 154(10), C562– C570.

[17] M. A. Khan, R. L. Williams and D. F. Williams, Biomaterials, 1999, 20(7), 631–637.

[18] C. V. Vidal and A. I. Muñoz, Corros. Sci., 2008, 50(7), 1954–1961.

[19] C. Bartels, J. Schultze, U. Stimming and M. Habib, Electrochim. Acta, 1982, 27(1), 129–140.

[20] S. D. Cramer, B. S. C. Jr and C. Moosbrugger: ‘ASM handbook volume 13B: Corrosion: Materials’, ASM International Materials Park, 2005.

[21] R. J. Narayan: ‘ASM Handbook, Volume 23 - Materials for Medical De- vices’, ASM international, 2012. REFERENCES 95

[22] Y. Yan, A. Neville, D. Dowson and S. Williams, Tribol. Int., 2006, 39(12), 1509–1517.

[23] M. Wimmer, J. Loos, R. Nassutt, M. Heitkemper and A. Fischer, Wear, 2001, 250(1-12), 129 – 139, 13th International Conference on Wear of Materials.

Paper C

Fretting Corrosion of Hafnium in Simulated Body Fluids

C.1. ABSTRACT 99

To be submitted

Fretting corrosion of Hafnium in Simulated Body Fluids

C.1 Abstract

Metals used in medical devices can be exposed to fretting in applications such as dental and joint implants. The effect of fretting, combined with the corrosiveness of body fluids, leads to enhanced degradation that can eventually result in failure of the implant. Hafnium has been suggested as an interesting material for biomedical applications due to its good biocompatibility and osteogenesis. However, its behaviour in fretting corrosion conditions has not been studied in depth. In order to investigate the effect of fretting on the corrosion resistance of hafnium and commercially pure (CP) titanium in simulated body fluids, a three electrode electrochemical cell integrated with a ball on flat reciprocating tribometer was used. The effect of fretting on the resistance of hafnium to localised corrosion was investigated. The susceptibility to localised corrosion was drastically decreased when hafnium was subjected to fretting. The study of the open circuit potential showed a more severe mechanical depassivation due to fretting in the case of CP titanium in comparison to hafnium. In addition, the anodic currents measured during potentiostatic tests were also higher in the case of CP titanium.

Keywords: Hafnium, Titanium, Biomaterial, Corrosion, Fretting corrosion 100 PAPER C. FRETTING CORROSION OF HAFNIUM

C.2 Introduction

The majority of metals used for biomedical applications owe their corrosion resistance to the formation of an oxide layer on their surface that isolates the bulk material from the corrosive media. In some applications, metals are exposed to small relative motion (fretting), which, combined with the corrosiveness of biological environments, leads to enhanced material degradation [1]. Mechanical damage could cause disruption of the passive layer which would leave the bulk material exposed to the corrosive media, leading to the release of metallic particles and ions into the body. The released degradation products can induce a number of biological responses such as inﬂammation, hypersensitivity reactions or allergy that can lead to the failure of the implant [2]. In addition, fretting-corrosion can accelerate crack initiation on the materials and lead to early failure of metallic devices [1].

Fretting corrosion can be found in fracture ﬁxation screws and plates, dental implants and joint implants [3, 4]. Titanium and titanium alloys are used for some of these applications due to their excellent corrosion resistance and biocompatibility. However, they show relatively poor wear characteristics that may lead to the release of degradation products in wear-corrosion conditions [5]. In spite of their good biocompatibility, titanium particles can induce the release of osteolytic cytokines involved in implant loosening [6, 7]. In addition, metallic particles can be disseminated in the body and accumulate in organs such as the liver and kidneys [8]. Research on the possibilities of new biocompatible materials that generate less degradation products could lead to a reduced risk of adverse biological reactions and improved outcomes.

Hafnium is a transition metal with high ductility and strength, and good resistance to corrosion and mechanical damage [9]. A very protective oxide layer, mainly formed by HfO2 [10], provides excellent corrosion resistance. Various studies have shown that hafnium has good biocompatibility and osteogenesis [11] and a tissue response similar to that observed for titanium [12]. However, there is limited information about the behaviour of hafnium in biological environments. Despite the passive state and the good corrosion behaviour of the material in simulated body ﬂuids, a tendency to suffer from localised corrosion has been reported in previous studies [13]. To date, the behaviour of hafnium in biological environments in fretting-corrosion conditions has not been studied in depth. In the present study, electrochemical techniques have been used in order to investigate the effect of fretting on the corrosion C.3. MATERIALS AND METHODS 101

Figure C.1: Schematic representation of the experimental set up. behaviour of hafnium, and to assess the potential of the material for use in biomedical applications.

C.3 Materials and Methods

C.3.1 Preparation of specimens and electrolyte solutions

Commercially pure titanium (Grade 2) and commercially pure hafnium rods (GfE Metalle und Materialien GmbH, Germany) were used in this study. Plates of 6 mm thickness and 25 mm diameter were prepared from the initial rod. Samples were polished with silicon carbide (SiC) paper up to 1200 grit followed by a 9 µm and a 3 µm diamond paste. Colloidal silica suspension was employed for the ﬁnal polishing. Samples were cleaned and rinsed with distilled water and ethanol before they were assembled into the measurement cell.

In order to simulate body ﬂuids, 25% bovine calf serum solution (Sigma- Aldrich, USA) was used. In addition, 0.9% NaCl solution was used in to isolate the effect of the organic species (proteins, amino acids, etc.) from the saline environment. All tests were conducted at 37◦C.

C.3.2 Microhardness measurements

A Matsuzawa MXT-CX microhardness tester (Matsuzawa Co., Japan) was used to study the microhardness of CP titanium and hafnium. At least ﬁve measurements were performed on each sample to ensure repeatability. 102 PAPER C. FRETTING CORROSION OF HAFNIUM

C.3.3 Fretting corrosion tests A ball on ﬂat reciprocating tribometer (CETR-UMT-2, Bruker Nano Inc., USA) was used to perform the fretting corrosion experiments. Samples were mounted on a polyoxymethylene holder and were fretted against a stationary silicon nitride ball (grade 10) of 12 mm diameter (Redhill Precision, Czech Reublic). A normal load of 10 N was used. The frequency was controlled at 1 Hz and the maximum stroke length was 120 µm. Tests were performed for 3600 cycles. A three electrode electrochemical cell was integrated into the system to perform the fretting corrosion tests. The schematic diagram of the fretting corrosion set up is shown in Figure C.1. The electrochemical cell consisted of a platinum wire used as the counter electrode, an Ag/AgCl electrode (3M KCl) used as the reference electrode, and the studied material was used as the working electrode. All potentials are given with respect to the Ag/AgCl electrode, which is 196 V on the standard hydrogen electrode scale. Three different electrochemical tests were performed in this study:

Cyclic Polarisation Tests Cyclic polarisation scans were performed on hafnium in fretting conditions in 0.9% NaCl and 25% serum solution. Before the cyclic polarisation scan, samples were immersed in each solution for 1800 s to stabilise them. There- after, fretting was started and the OCP was measured until a stable value was reached. Samples were polarised in the cathodic direction down to -1 V at a scan rate of 10 mV s−1. Finally, the cyclic polarisation test was performed at a scan rate of 2 mV s−1 up to a potential of 3 V or until the current density limit (800 µAcm−2) was reached. Tests were repeated at least three times.

Open Circuit Potential (OCP) Tests The OCP of the studied materials was measured as a function of time. The test consisted of three different phases. First, the sample was immersed in static conditions for 3600 s. Then, the speciﬁc load was applied and the fretting was started. Finally, after 3600 cycles fretting was stopped and the sample was unloaded. Tests were repeated at least three times.

Potentiostatic Tests A constant passive potential of 0 V (vs Ag/AgCl) was applied to monitor the current as a function of time and the effect of the mechanical depassivation due to fretting. The tests consisted of three phases: ﬁrst, the current was monitored C.4. RESULTS 103

Table C.1: Microhardness (HV500g) of hafnium and CP titanium. HV500g Hf 213 ± 3 CP Ti 163 ± 6 ±: Standard deviation

Table C.2: Comparison of the breakdown potential of hafnium in fretting con- f retting static ditions (Eb ) and in static conditions (Eb ) in 0.9% NaCl and 25% serum. f retting static Eb (V) Eb (V) [13] 0.9% NaCl 0.279 ± 0.003 1.387 ± 0.080 25% serum 0.405 ± 0.007 2.331 ± 0.128 ±: Standard deviation at 0 V during 3600 s before the start of fretting; after that, the sample was fretted during 3600 s; ﬁnally, fretting was stopped and currents were monitored for 3600 s. Tests were repeated at least three times.

C.4 Results

Microhardness measurements Table C.1 shows the microhardness values measured for each material. Hafnium showed higher microhardness in comparison to CP titanium.

Cyclic Polarisation measurements The potentiodynamic scans of hafnium in 0.9% NaCl and 25% serum solutions (Figure C.2) showed a breakdown of the passive layer due to the formation of pits on the surface. The formation of the pits and the subsequent localised corrosion were retarded in 25% serum (Table C.2). Once the potential scan was reversed, the pit propagation caused that the currents continued increasing. The currents reached in the saline solution were higher than in the serum solution. The repassivation potential, at which currents reached the original passive values, was slightly higher in the case of 25% serum solution.

Open Circuit Potential (OCP) measurements The OCP measurements are shown in Figure C.3. The tests can be divided in 104 PAPER C. FRETTING CORROSION OF HAFNIUM

1.00

0.75

0.50

0.25

0.00 -0.0002 0 0.0002 0.0004 0.0006 0.0008 0.001 0.0012

-0.25

Potential (V) vs Ag/AgCl Potential (V) vs -0.50 Hf - 0.9% NaCl -0.75 Hf - 25% serum -1.00 Current Density (A cm-2)

Figure C.2: Cyclic polarisation of hafnium in fretting conditions in 0.9% NaCl and 25% serum.

Table C.3: Maximum drop (ΔEmax) and average variation (ΔEavg) of the OCP due to mechanical depassivation caused by fretting. ΔEmax (mV) ΔEavg (mV) 0.9% NaCl 79.2 ± 25.9 36.9 ± 1.6 Hf 25% serum 151.9 ± 36.3 39.9 ± 19.7 0.9% NaCl 227.3 ± 43.6 81.3 ± 5.5 CP Ti 25% serum 260.4 ± 22.0 171.9 ± 21.3 ±: Standard deviation

(a) (b)

Figure C.3: Effect of fretting on the open circuit potential of hafnium and CP titanium in (a) 0.9% NaCl and (b) 25% serum. C.4. RESULTS 105

(a) (b)

Figure C.4: Evolution of the current as a function of time at a potential of 0V in (a) 0.9% NaCl and (b) 25% serum. three phases: from A to B (in Figure C.3), before fretting started, an anodic shift was observed which suggested a thickening of the passive layer; when fretting started at point B (in Figure C.3), the OCP values dropped towards more active values. The OCP reached a minimum value immediately after the onset and it tended to increase towards less active values due to the equilibrium between the mechanical depassivation and electrochemical repassivation. Once the fretting was stopped at point D (in Figure C.3), the OCP recovered towards more positive values indicating repassivation of the material. The values of the maximum variation of the OCP and the average variation during fretting are shown in Table C.3. Despite titanium showed more noble OCP values in static conditions, the variation of the values due to fretting was larger than in the case of hafnium in the two studied solutions. The average drop in the OCP (ΔEavg) due to fretting was found higher for CP titanium, being more pronounced in 25% serum than in the saline solution. The inital depassivation peak (ΔEmax) was more acute in the case of CP titanium in both solutions. Furthermore, the depassivation peak was more pronounced in 25% serum in hafnium and CP titanium.

Potentiostatic tests The evolution of the current measured in CP titanium and hafnium at a constant passive potential (0 V vs Ag/AgCl) is shown in Figure C.4. The test can be divided in three phases: ﬁrst from A to B (in Figure C.4), samples were immersed in the solution and a passive potential (0 V vs Ag/AgCl) was applied. The low currents indicated that both materials were in a passive state. After that, from B to C (in Figure C.4), the onset of the fretting led to removal of the 106 PAPER C. FRETTING CORROSION OF HAFNIUM

Table C.4: Anodic peak current (Ipeak) and average current (Iavg) measured during fretting at a potential of 0 V (vs Ag/AgCl) for Hf and CP Ti. Ipeak (µA) Iavg (µA) 0.9% NaCl 6.01 ± 3.24 0.86 ± 0.25 Hf 25% serum 4.35 ± 2.15 1.03 ± 0.22 0.9% NaCl 14.42 ± 6.16 2.82 ± 0.55 CP Ti 25% serum 23.67 ± 6.35 3.57 ± 0.52 ±: Standard deviation passive ﬁlm, causing a sudden increase in the current due to the exposition of the unprotected metal to the media. After an initial peak, the current decreased towards lower values due to the equilibrium between mechanical repassivation and electrochemical depassivation. Finally, from C to D (in Figure C.4), the low currents were quickly recovered when the mechanical damaging stopped, indicating that the passive layer had reformed.

The peak current and the average current measured are summarised in Table C.4. The peak currents measured in the case of CP titanium were higher compared to hafnium, indicating a more severe initial depassivation. The average currents were higher in the case of CP titanium as well; however, Figure C.1(a) shows that initially the currents measured were similar in both materials, increasing after this initial stage in the case of CP titanium. This could have been caused by a higher wear and an increased depassivated area.

C.5 Discussion

The behaviour of hafnium and CP titanium in fretting corrosion conditions has been studied in order to assess the suitability of hafnium for biomedical applications.

Previous studies have shown a tendency of hafnium to suffer from localised corrosion in simulated body ﬂuids [13]. The effect of fretting on the breakdown potential of hafnium was studied by means of cyclic polarisation scans. The breakdown potential observed when hafnium was subjected to fretting was drastically reduced in comparison to the values observed in static conditions (Table C.2). Breakdown potential was higher in the serum solution, which C.5. DISCUSSION 107 suggested that the presence of organic species retarded the formation of corrosion pits and increased the repassivation potential when the pits were formed. This could be due to the negative charge of proteins at the pH of the studied solutions (pH 7.4) [14], which would lead to attractive forces towards the anodically polarised surface. Negatively charged proteins and chlorides, which are known to be a cause of localised corrosion [15], would compete to adsorb on the surface, retarding the formation of pits [16]. The currents reached due to the propagation of the pits were reduced when proteins were present in the media. A similar effect of the organic species has been reported in static conditions in previous studies [13].

Open circuit potential (OCP) is commonly used as a qualitative indication of the corrosion regime in which a material resides [17]. Both CP titanium and hafnium were in the passive state when immersed in saline and serum solutions. Both materials showed an anodic shift before the onset of the fretting that suggested thickening of the passive ﬁlm. When fretting started a sudden cathodic shift was observed. The mechanical damage on the surface exposed the bulk metal to the corrosiveness of the media; therefore the OCP value was much closer to that of the fresh material [18]. The OCP values measured cor- respond to the mixed potential between the fretted and the unfretted area. An initial peak in the cathodic direction was observed with the onset of the fretting. The initial peak was more prominent in the case of CP titanium, which suggested a more severe mechanical depassivation when fretting started. After the initial cathodic peak the OCP increased towards more anodic values due to the balance between the mechanical depassivation due to fretting and the electrochemical repassivation or reformation of the oxide layer due to the reaction of the materials with the oxygen in the solution [19]. The difference between the OCP value on the passive state and under fretting was smaller in the case of hafnium, suggesting less mechanical damage, which could be related to the higher hardness of the material. A similar trend was observed in both the saline solution and the serum solution.

The evolution of the anodic current during the tribological tests under passive potential (0 V vs Ag/AgCl) showed higher corrosion currents for CP titanium under the studied conditions. Current increased drastically when the mechanical damage started. The initial peak was higher for CP titanium in both 0.9% NaCl and 25% serum solutions, which could be caused by a more severe initial depassivation due to the lower hardness of the material. After the initial peak, lower and more stable current values were reached. The average 108 PAPER C. FRETTING CORROSION OF HAFNIUM currents measured for hafnium were lower than those for CP titanium in the saline solution and the 25% serum solution. In the case of hafnium in 0.9% NaCl solution, the currents remained relatively low during the time of the test, however, in the case of hafnium in 25% serum and CP titanium in both solutions, an increase of the current was observed during the test. Proteins could have had an effect on the passivation of the materials by restricting the diffusion of oxygen to the surface and making repassivation of the metal more difﬁcult [14]. The negative effect of proteins on the repassivation of the metal would increase not only the corrosion but also the wear induced by corrosion, which could cause an increase in total wear-corrosion damage. In the case of CP titanium, the lower surface hardness might lead to increased mechanical damage and therefore more severe mechanical damage.

C.6 Conclusions

Metals used for medical implants can be subjected to fretting corrosion, which can compromise the life time of the implant. The fretting corrosion behaviour of hafnium and CP titanium has been investigated. The main conclusions of this study are: • Hafnium tended to suffer from localised corrosion; its resistance was drastically decreased when the material was subjected to fretting. How- ever, the organic species contained in the serum solution retarded the formation of pits and limited their propagation. • Hafnium showed a passive state in both 0.9% NaCl and 25% serum solution under static conditions. The passive layer was damaged due to fretting; however, the mechanical depassivation was more severe in the case of CP titanium in comparison to hafnium. Both hafnium and CP titanium repassivated in the studied solutions when the mechanical damaging stopped. • The damage produced due to fretting produced increased currents and therefore higher corrosion of the materials. Hafnium showed lower anodic currents in fretting than CP titanium.

C.7 Acknowledgements

This work was funded by the Centrum för Medicinsk Teknik och Physic (CMTF) and Kempe Stiftelserna. REFERENCES 109

References

[1] S. Hiromoto and S. Mischler, Wear, 2006, 261(9), 1002–1011.

[2] N. J. Hallab and J. J. Jacobs, Bull. NYU Hosp. Jt. Dis., 2009, 67(2), 182– 188.

[3] D. Upadhyay, M. A. Panchal, R. Dubey and V. Srivastava, Mater. Sci. Eng., A, 2006, 432(1), 1–11.

[4] J. Jacobs, J. Gilbert and R. Urban, J. Bone Joint Surg. Am., 1998, 80A(2), 268–282.

[5] M. Geetha, A. Singh, R. Asokamani and A. Gogia, Progress in Materials Science, 2009, 54(3), 397–425.

[6] N. Taki, J. M. Tatro, J. L. Nalepka, D. Togawa, V. M. Goldberg, C. M. Rimnac and E. M. Greenﬁeld, J. Orthop. Res., 2005, 23(2), 376–383.

[7] S. Suñer, J. Tipper and N. Emami, Tribol. - Mater. Surf. Interfaces, 2012, 6(2), 39–52.

[8] R. M. Urban, M. J. Tomlinson, D. J. Hall and J. J. Jacobs, The Journal of arthroplasty, 2004, 19(8), 94–101.

[9] E. Cerreta, C. Yablinsky, G. Gray, S. Vogel and D. Brown, Mater. Sci. Eng., A, 2007, 456(1), 243–251.

[10] W. A. Badawy and F. M. Al-Kharaﬁ, J. Mater. Sci., 1999, 34, 2483–2491.

[11] H. Matsuno, A. Yokoyama, F. Watari, M. Uo and T. Kawasaki, Biomate- rials, 2001, 22(11), 1253–1262.

[12] S. Mohammadi, M. Esposito, M. Cucu, L. E. Ericson and P. Thomsen, J. Mater. Sci.: Mater. Med., 2001, 12(7), 603–611.

[13] J. Rituerto Sin, A. Neville and N. Emami: ‘3rd International Tribology Symposium of IFToMM, Luleå, Sweden’.

[14] M. A. Khan, R. L. Williams and D. F. Williams, Biomaterials, 1999, 20(7), 631–637.

[15] D. D. Macdonald, J. Electrochem. Soc., 1992, 139(12), 3434–3449.

[16] C. V. Vidal and A. I. Muñoz, Corros. Sci., 2008, 50(7), 1954–1961. 110 REFERENCES

[17] Y. Yan, A. Neville and D. Dowson, Wear, 2007, 263(7), 1105–1111.

[18] B. Sivakumar, S. Kumar and T. Narayanan, Tribol. - Mater. Surf. Inter- faces, 2011, 5(4), 158–164.

[19] M. Dimah, F. D. Albeza, V. A. Borrás and A. I. Muñoz, Wear, 2012.