Investigation of Hafnium for Biomedical Applications

LICENTIATE T H E S I S

Department of Engineering Scienceand MathematicsDivision of Machine Elements

Investigation of Hafnium for Biomedical Applications

Corrosion and Tribocorrosion in Simulated Body Fluids

Jorge Rituerto Sin

ISSN: 1402-1757ISBN 978-91-7439-680-5 (print)ISBN 978-91-7439-681-2 (pdf)

Luleå University of Technology 2013

Jorge Rituerto Sin Investigation of H

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ISSN: 1402-1757 ISBN 978-91-7439-XXX-X Se i listan och fyll i siffror där kryssen är

INVESTIGATION OF HAFNIUMFOR BIOMEDICAL APPLICATIONS

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Corrosion and Tribocorrosionin Simulated Body Fluids

JORGE RITUERTO SIN

Luleå University of TechnologyDepartment of Engineering Science and Mathematics

Division of Machine Elements

Printed by Universitetstryckeriet, Luleå 2013

ISSN: 1402-1757 ISBN 978-91-7439-680-5 (print)ISBN 978-91-7439-681-2 (pdf)

Luleå 2013

www.ltu.se

Cover figure: Schematic representation ofan atom of hafnium.

Title page figure: Energy dispersive X-Ray spectroscopy(EDX) spectrum of hafnium.

INVESTIGATION OF HAFNIUMFOR BIOMEDICAL APPLICATIONS

Corrosion and Tribocorrosionin Simulated Body Fluids

Copyright c© Jorge Rituerto Sin (2013).This document is freely available at

http://www.ltu.se

or by contacting Jorge Rituerto Sin,

[email protected]

This document may be freely distributed in its original form including the cur-rent author’s name. None of the content may be changed or excluded withoutthe permission of the author.

Printed by Universitetstryckeriet, Luleå 2013

ISSN:ISBN:

Luleå 2013

Preface

The work presented in this Licentiate thesis has been carried out at the Divisionof Machine Elements at Luleå University of Technology (Luleå, Sweden) andat 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) andKempe Stiftelserna. The Swedish Research School in Tribology is acknowl-edged 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 AnneNeville, 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 laband Lars Frisk for valuable help with sample preparation.

I am grateful to my friends and colleagues at Leeds University, especiallyJames Hesketh for long discussions about tribocorrosion and for his help inthe 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 afriendly and enjoyable place to work. I would especially like to express mygratitude to Gregory Simmons for his valuable feedback on my manuscriptsand Silvia Suñer for all her support and advice that help me to improve everyday.

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 tough-ness, which make them the material of choice for many biomedical applica-tions. However, the main drawback of metals is their general tendency to cor-rode, which is an important factor when they are used as biomaterials due tothe corrosive nature of the human body.

Titanium and titanium alloys are widely used in biomedical devices due totheir excellent corrosion resistance and good biocompatibility. However, oneof the disadvantages of titanium is its low wear resistance.

Hafnium is a passive metal with a number of interesting properties, such ashigh ductility and strength, as well as resistance to corrosion and mechanicaldamage. Previous studies have shown that hafnium has good biocompatibilityand osteogenesis. However, the behaviour of hafnium in biological environ-ment has not been studied in great depth. Furthermore, little is known aboutthe resistance of the passive layer under wear-corrosion conditions and the ef-fect of proteins on its corrosion and tribocorrosion behaviour.

The overall goal of this study is to assess the potential of hafnium for use inbiomedical applications. The aim of this work is to investigate the corrosionresistance of hafnium in simulated body fluids as well as its behaviour in wearcorrosion and fretting corrosion conditions.

The results showed that hafnium presents a passive state in the presence ofproteins and its oxide layer provides high protection to corrosion. In addi-tion, although the passive layer could be disrupted due to wear and fretting,increasing the corrosion of the metal, it was rapidly rebuilt when the damagingceased. On the other hand, the main drawback of hafnium was its tendency tosuffer from localised corrosion. Although the formation of corrosion pits wasretarded in the presence of proteins, it was drastically increased when hafniumwas scratched or subjected to fretting.

Appended Papers

[A] J. Rituerto Sin, X. Hu, N. Emami. Tribology, Corrosion andTribocorrosion 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, September18-21, 2011, London, UK.∗Reprinted with permission of Maney Publishing and Tribology - Ma-terials, Surfaces and Interfaces. Permission to reprint this material isgratefully acknowledged.

[B] J. Rituerto Sin, A. Neville, N. Emami. Corrosion and Tribocor-rosion of Hafnium in Simulated Body Fluids. Submitted forjournal 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 ofHafnium 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 31.1 The need for biomaterials . . . . . . . . . . . . . . . . . . . . 3

1.1.1 Types of biomaterials . . . . . . . . . . . . . . . . . . 41.2 Metallic biomaterials . . . . . . . . . . . . . . . . . . . . . . 5

1.2.1 Titanium and titanium alloys . . . . . . . . . . . . . . 71.2.2 Possibilities of hafnium . . . . . . . . . . . . . . . . 8

1.3 Tribocorrosion of metallic implants . . . . . . . . . . . . . . 101.3.1 Corrosion of implant materials . . . . . . . . . . . . . 101.3.2 Wear of implant materials . . . . . . . . . . . . . . . 111.3.3 Case study: Tribocorrosion of Metal on Metal implants 12

2 Objectives 15

3 Materials and Methods 173.1 Materials and material preparation . . . . . . . . . . . . . . . 173.2 Test conditions . . . . . . . . . . . . . . . . . . . . . . . . . 173.3 Experimental set up . . . . . . . . . . . . . . . . . . . . . . . 18

3.3.1 Electrochemical cell . . . . . . . . . . . . . . . . . . 183.3.2 Tribocorrosion set up . . . . . . . . . . . . . . . . . . 18

3.4 Experimental methods . . . . . . . . . . . . . . . . . . . . . 193.4.1 Electrochemical tests . . . . . . . . . . . . . . . . . . 193.4.2 Tribocorrosion tests . . . . . . . . . . . . . . . . . . . 203.4.3 Fretting corrosion tests . . . . . . . . . . . . . . . . . 21

3.5 Surface analysis . . . . . . . . . . . . . . . . . . . . . . . . . 223.5.1 Microhardness . . . . . . . . . . . . . . . . . . . . . 223.5.2 Optical microscopy . . . . . . . . . . . . . . . . . . . 223.5.3 3D optical profilometry . . . . . . . . . . . . . . . . . 23

vii

4 Results and Discussion 254.1 Corrosion resistance of hafnium . . . . . . . . . . . . . . . . 254.2 Tribocorrosion of hafnium . . . . . . . . . . . . . . . . . . . 284.3 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

1

Chapter 1

Introduction

In this chapter, the need for biomaterials is introduced. The most commonlyused biomaterials are summarised, giving especial attention to the main groupsof metals used for biomedical applications. Titanium and titanium alloys arediscussed in more detail, and the possibilities of hafnium as an implant ma-terial are explored. Finally, the environment and conditions that biomaterialsface 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 oneunit. When one or more of these organs are damaged due to injuries, infectionsor different types of diseases, body function is affected. Different solutionshave been developed to minimise the adverse effect of these injuries and dis-eases. Often, these solutions involve the implantation of biomedical devices inthe 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 apart or a function of the body in a safe, reliable, economic, and physiologicallyacceptable manner [1]. The implantation of biomaterials in the body impliesdifferent restrictions that must be considered to minimise the risk of failure andthe need of revision of the implant (Figure 1.1). A biomedical device shouldnot 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 implantmust 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 surgeryoften combine biological aspects with material and engineering aspects.

ical properties of the material must be adequate for the requirements of eachparticular application. The machining properties, the cost and the availabilityare 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 propertiescan be normally improved by alloying and by thermal and mechanicaltreatments. However most of the metals are highly reactive and thereforethey can corrode unless they are protected. Screws for dental implantsand fracture fixation devices are some of their applications.

• Ceramics: Ceramics are stiff, hard, biocompatible and resistant to cor-rosion. However, their brittleness can be a concern when used for someapplications, such as load bearing implants. The crown of dental im-plants and the femoral head of certain types of hip implants can be madeof 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 theareas of application.

• Composites: composites combine two or more materials to obtain a ma-terial with improved characteristics. Their properties can be adaptedaccording to certain requirements, however they can be difficult to pro-duce and they are normally expensive. Composites are used for differentapplications such as spine cages and acetabular cups for hip implants.

• Natural biomaterials: Beside the groups of synthetic materials, materi-als derived from animals or plants can be grouped as natural biomateri-als. 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. Theycan denature and decompose more easily than most synthetic materialsand they may have problems with immunogenicity. One example of thisgroup is collagen, which has been used in tissue engineering for appli-cations such as skin replacement, bone substitutes, and artificial bloodvessels 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 or-der to select the best candidate. The behaviour of the candidate material underthe specific conditions must be understood in order to improve the outcomes ofthe biomedical implants. The field of biomaterials is highly interdisciplinaryand requires collaboration between clinicians, biologists, materials scientistsand engineers in order to develop new biomaterials and improved medical de-vices.

1.2 Metallic biomaterials

Metals are generally defined as polycrystalline compounds formed by metal-lic bonding, where an electron cloud surrounds positively charged metal ions(Figure 1.2) [3]. Metals present excellent mechanical properties such as highstrength, ductility and toughness, which make them the materials of choice forvarious biomedical applications. Among the multiple uses of metals, cranialplates, dental implants, fracture fixation plates and screws, joint implants arecommon. Metallic biomaterials are also used for various devices such as cagesfor pumps, valves and heart pacemakers, where materials do not support highloads. Various types of surgical instruments are also made of metal.


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 istheir tendency to corrode. Most of the metals used for biomedical applicationsare passive metals, an oxide film is spontaneously formed on their surfacewhich protects the material from the corrosive media.

In the 1900s, a metal alloy known as "vanadium steel" was specifically de-signed to be used for temporary fixation plates [4]. Over the years, vanadiumsteel was replaced with more resistant materials. There are three main groupsof metals that are used in medical applications: stainless steels, cobalt basedalloys 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 isthe 316L. The presence of chromium provides the passive state to the al-loy. However, stainless steel tends to corrode under stresses and in oxy-gen depleted regions. Therefore, it is susceptible to suffer from creviceand pitting corrosion [5]. For this reason, stainless steel is normally onlyused in temporary devices such as fracture fixation plates and screws, aswell as for surgical instruments.


• Cobalt-based alloys. The low wear resistance of stainless steel led tothe development of a cobalt-chromium-molybdenum alloy known as Vi-tallium [6]. Today there are two main categories of cobalt-based al-loys, Cobalt-Nickel-Chromium-Molybdenum (CoNiCrMo) alloys andCobalt-Chromium-Molybdenum (CoCrMo) alloys, however, CoCrMoalloys are the most commonly used. The presence of chromium asan alloying element gives these alloys excellent corrosion resistance,which, together with their good wear resistance and mechanical prop-erties, makes them suitable for applications where the device is subjectto 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 excel-lent corrosion properties and biocompatibility for static implant condi-tions. However, their main disadvantage is their poor wear resistance.In addition to comercially pure titanium, the most commonly used ti-tanium alloy is titanium-6% aluminium-4% vanadium (Ti-6Al-4V) [7].Titanium and titanium alloys are used for many biomedical applicationssuch as bone screws, dental screws, pacemaker cases, femoral stems inhip 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 designeda 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 withtitanium (which he named as osseointegration) [10], titanium started to be ex-tensively used in dental and joint replacement applications. Today, more than1.000 tonnes of titanium are implanted into patients in the form of biomedicaldevices in the world every year [11].

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


Figure 1.3: Schematic representation of the reactions of the ions released intothe 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 andthe tissue response (Figure 1.3). Ions can bind with biomolecules formingcomplexes that can activate the immune system. More active ions will reactmore easily with hydroxyl radicals and anions forming oxides and salts, de-creasing the possibility of reacting with biomolecules. Less active ions willstay in the system longer and, subsequently, they will have more opportunitiesto react with biomolecules and reveal toxicity [13]. Therefore, the high reac-tivity of titanium not only provides it with good corrosion resistance, but alsowith low toxicity.

In spite of the good corrosion properties of titanium, it provides relativelypoor wear characteristics [7]. Although titanium shows a low toxicity, therelease of particles has been related to the release of osteolitic cytokines thateventually lead to implant loosening [14, 15]. Furthermore, they can be spreadinto 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 withtitanium and zirconium. The metal was first identified by Dirck Coster and


Georges de Hevesy in Copenhagen in 1923 [17] and owes its name to ’Hafnia’,the Latin name for Copenhagen. Hafnium is always found associated withzirconium in ores. The main mineral where it is found is zircon, with a ratioHf/Zr of about 2.5% [18]. Due to a number of interesting properties such ashigh ductility and strength, as well as resistance to corrosion and mechanicaldamage [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 insome superalloys used in aircrafts engines [20].

In 1984, Marcel Pourbaix included hafnium among the group of metals to beconsidered 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 bodyat that time it was discarded from the final list of metals to be theoreticallyconsidered [21]. The passive layer found on hafnium’s surface in aqueoussolutions is mainly formed of HfO2, according to the equation (1.2). Thisoxide film is highly stable in neutral solutions [22].

H f +2H2O↔ H fO2+4H++4e− (1.2)

More recently, the properties of hafnium as an implant material have beeninvestigated [23, 24]. Matsuno et al. [23] implanted a number of refractorymaterials (titanium, hafnium, niobium, tantalum and rhenium) in subcutaneoustissue and femoral bone marrow of rats, and performed histological observa-tion and elemental mapping in order to evaluate the biocompatibility and os-teogenesis of such materials. The results of the in vivo study showed thathafnium had good biocompatibility and osteogenesis. A different study on thetissue response to hafnium [24] showed a similar response in soft and hardtissues in two different animal species (rat and rabbit), which suggests thathafnium might be an interesting material for biomedical applications.

Hafnium has also been investigated as an alloying element in titanium al-loys. Different Ti-Hf binary alloys have been reported in the literature. WroughtTi-Hf alloys have shown a gradual increase in yield and tensile strength withincreased amount of hafnium, up to 40%, whereas the ductility decreased [25].More recent studies also showed an increased yield and tensile strength as wellas bulk hardness for cast Ti-Hf alloys [26]. The electrochemical behaviour ofthis type of alloys is similar to pure titanium, showing an excellent corrosionresistance [27]. Ti-Nb-Hf-Zr alloys have also been investigated [28]. Thesealloys have shown a low elastic modulus which is beneficial in order to reduce


the stress shielding effect and to enhance bone growth. It has also been shownthat 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 notbeen studied in depth. Little is known about the resistance of the passive layerto wear-corrosion conditions and the effect of proteins on its corrosion andtribocorrosion behaviour. Thus, further studies of hafnium under biologicalconditions are needed in order to determine the suitability of this material forbiomedical 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 withbody fluids. The chloride ions dissolved in body fluids make them a corrosiveenvironment for metallic materials. In addition, different organic species suchas amino acids and proteins are also found in biological fluids, which influencethe corrosion of metals [30]. Changes in the pH can also affect corrosion. ThepH of the body normally remains around 7.4 because body fluids are bufferedsolutions [31]. The implantation of a device into a hard tissue can lead to pHvalues of around 5.6 or in the case of infection, the pH can reach values ofaround 9 [21].

Corrosion can be defined as the degradation of a material caused by the re-action with its surrounding environment. The corrosion phenomena in humanfluids is due to an electrochemical processes, which implies two reactions: theoxidation of a metal, or anodic reaction (Eq. 1.3); and the reduction of anoxidising 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 themain types of corrosion are:

• General corrosion, also called uniform corrosion, affects the entire sur-face of the material, leading to the dissolution of metallic ions in themedia.

• Pitting corrosion is a type of localised corrosion that consists of theformation of small cavities at the surface of the metal. The oxidationreaction in the pit leads to lower pH and the migration of chloride ionsto balance the charge of the metal ions. This creates a more aggressiveenvironment locally and enhances the corrosion, penetrating at a highrate.

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

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

• Fretting corrosion takes place at the interface of contacting surfaceswith small relative motion between them. It leads to an enhanced corro-sion 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 butalso to variable and complex service conditions. The static and dynamic loadsaffecting medical devices greatly depend on the activity of the patient. Inaddition, the application of the device will ultimately determine the type ofconditions that will affect its degradation. Although passive state of metallicbiomaterials minimises the failure rate directly caused by corrosion, differenttypes of wear can damage the oxide layer and lead to accelerated corrosion


Figure 1.4: Combinde effect of wear and corrosion in metal on metal hip im-plants.

processes. The bearing surfaces of joint implants are a clear example of theoccurrence of wear in medical devices.

Deterioration by small relative motion between solids, normally known asfretting, can also be found in medical devices such as fracture fixation screwsand plates, at the head/neck interface of modular hip implants or at the ce-ment/stem interface of hip implants. Fretting leads to mechanical damage ofthe oxide layer and wear of the metallic surfaces. The oxidised metallic par-ticles also contribute to accelerated wear. The interaction between the me-chanical contact, the electrochemical processes and the surrounding biologicalenvironment determines the material loss and the generation of metallic parti-cles.

1.3.3 Case study: Tribocorrosion of Metal on Metal implants

Metal on metal (MoM) hip replacements have been considered as an alterna-tive 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 thedevice. MoM implants (Figure 1.4) consist of a metallic stem that is insertedinto the femur, a metallic femoral head and a metallic acetabular cup that ar-ticulates with the head. The excellent wear properties of cobalt chromiummolybdenum alloys makes it the material of choice for the bearing surfaces ofthe implant. It has been reported that MoM implants generate approximately100-fold less volumetric wear than MoP [32] and the particle size has beenestimated in the range of 60 to 90 nm [33], which ideally avoids the host re-sponse produced in the case of MoP. However, the mechanical damage causedon the bearing surface due to wear increases the corrosion of the implants andthe release of metallic ions. Higher ion levels have been related with variousbiological reactions such as hypersensitivity, pseudotumours and carcinogene-sis [34, 35, 36], which eventually lead to failure of the implant. Recent studieshave also shown the importance of metallic ions generated in the contact be-tween the cement and the femoral stem [37] and on the head and neck contactof modular implants due to fretting corrosion [38]. Today, MoM implants arethe subject of controversy due to the concerns related the biological responseto 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 materialfor medical devices. Hafnium has shown good biocompatibility and osteogen-esis, comparable to that observed for titanium. Therefore it has been suggestedas an interesting material for biomedical applications [24].

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

The specific objectives of this research work were:

• To understand and summarise the importance of tribocorrosion in thefield of biomaterials. The particular case of metal on metal implants wasstudied 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 insimulated body fluids.

15

Chapter 3

Materials and Methods

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

3.1 Materials and material preparation

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

3.2 Test conditions

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

17

18 CHAPTER 3. MATERIALS ANDMETHODS

Figure 3.1: Schematic representation of the three electrode electrochemicalcell used.

3.3 Experimental set up

The main goal of corrosion testing is to study the corrosion parameters of im-plant materials and to predict their behaviour when implanted into the body.Furthermore, new materials developed for medical applications must be stud-ied not only in terms of corrosion resistance but the effect of wear and frettingshould also be considered since they will affect the performance of the mate-rial.

3.3.1 Electrochemical cell

In order to study the corrosion behaviour of CP titanium and hafnium, a threeelectrode electrochemical cell was used (Figure 3.1). The electrochemical cellconsisted of a platinum wire, that was used as the counter electrode; a Ag/AgClelectrode (3 M KCl), used as the reference electrode (196 mV vs SCE); andthe studied specimens, that acted as the working electrode. All electrochemicalmeasurements 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 tri-bological tester (Figure 3.2). The configuration of the electrochemical cell wasthe same as the previously detailed.

3.4. EXPERIMENTALMETHODS 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 weartester that conforms to ASTM G133. Specimens were mounted on a poly-oxymethylene holder and were rubbed against a silicon nitride ball of 12 mmin diameter. A normal load of 40 N was applied. The frequency was controlledat 1 Hz and the stroke length was 10 mm. The specimens were passivated in30% HNO3 for 30 minutes before each test according to the standard ASTMF-86. A digital microbalance with 0.01 mg resolution was used to gravimetri-cally 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 againsta silicon nitride ball of 12 mm in diameter. A normal load of 10 N was appliedand the frequency was controlled at 1 Hz. A maximum stroke length of 120µm 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 amaterial to suffer from localised corrosion [31], which is determined by the


potential at which the anodic current rapidly increases. In addition, the rate ofmetal dissolution over a range of potentials can be studied. In order to carry outthe cyclic polarisation tests, specimens were mounted in the electrochemicalcell after polishing. The tests consisted of the following steps:

• The open circuit potential was measured for 3600 s to stabilise the ma-terial surface.

• The specimen was cathodically polarised at a scan rate of 10 mV s−1down 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 µAcm−2) was reached.

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

3.4.2 Tribocorrosion tests

Open circuit potential (OCP) testsAt the OCP, the net current is zero because the anodic reaction rate is equal tothe cathodic reaction rate. The OCP gives a qualitative indication of the stateof a metal (active or passive) in a particular environment (Figure 3.3) [39].The OCP can be monitored while the materials are subjected to mechanicaldamage. A drop in the OCP indicates a more active surface, which suggeststhat the passive layer is being damaged. The OCP was monitored during thethree 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 for3600 s to study the repassivation of the materials.

The percentage of repassivation of the OCP after mechanical depassivation hadended was evaluated according to Eq. 3.1:

Repassivation (%) =Et −Ewear

Epassive−Ewear (3.1)

Where Et is the OCP value at a time t; Ewear is the OCP value during wearbefore the mechanical damage ended and Epassive is the OCP value in staticconditions before the wear test starts.

3.4. EXPERIMENTALMETHODS 21

Pot

entia

l (V

)

Active material

Loading Unloading

Time (s)

Passive material

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

3.4.3 Fretting corrosion tests


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

• The open circuit potential was measured for 1800 s to stabilise the ma-terial surface.

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

• The specimen was cathodically polarised at a scan rate of 10 mV s−1down 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 µAcm−2) was reached.

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

• Fretting was stopped.


Open circuit potential (OCP) tests

The effect of fretting on the OCP was studied. The OCP was monitored duringthree 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 themetal is applied and the anodic current is monitored. The mechanical depassi-vation produced by fretting and its effect on the current at a given potential canbe analysed. A passive potential of 0 V (vs Ag/AgCl) was applied. The anodiccurrent was measured during the test, which consisted on the following steps:

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

• Fretting was started and the anodic current was measured for 3600 s tostudy 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 thepenetration of another body. The microhardness of CP titanium and hafniumwas studied with a microhardness tester using a load of 4903 mN according tothe ASTM E-384.

3.5.2 Optical microscopy

Optical microscopy was used to study the material surface after corrosion test-ing. 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 mea-surements of the surface topography (Figure 3.4). A Wyko 1100NT opticalinterferometer was used to study the cross section profile of the wear tracksafter tribocorrosion tests.

Chapter 4

Results and Discussion

This chapter summarises the findings of this work. The corrosion resistanceof hafnium is evaluated and compared to that of CP titanium, and the effect ofwear 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 studiedin 0.9% NaCl solution and 25% serum solution. Cyclic polarisation scans wereperformed to investigate the susceptibility to localised corrosion and the rateof anodic dissolution over a range of potentials. The cyclic polarisation curvesof hafnium in 0.9% NaCl and 25% serum are shown in Figure 4.1. A rapidincrease of the current density was observed in both solutions, which was dueto the breakdown of the passive layer and the formation of corrosion pits on thesurface. Pits were observed under optical microscopy in 0.9% NaCl and 25%serum solution (Figure 4.2 (a) and (b)). The breakdown potential (Eb) washigher in the serum solution. Furthermore, the repassivation potential (Er), atwhich current density reaches passive values after pits are formed, was alsoincreased by the presence of organic species in the solution. This could bedue to the anodic polarisation of the specimens during the tests, which wouldattract the negative charged proteins [30] as well as the chlorides dissolved inthe media. Chloride ions are known to be a cause of pit initiation, therefore,the competition between proteins and chlorides to adsorb on the surface wouldretard the formation of the pits [40].When the current limit (500 µA cm−1) was reached, the potential slope was re-versed. However, the current density continued increasing when the potential

25

26 CHAPTER 4. RESULTS AND DISCUSSION

-1.0

-0.5

0.0

0.5

1.0

1.5

2.0

2.5

3.0

0.0E+00 5.0E-04 1.0E-03 1.5E-03 2.0E-03 2.5E-03 3.0E-03

Pote

ntia

l (V)

vs

Ag/

AgC

l

Current Density (A cm-2)

Hf - 0.9% NaClHf - 25% serum

-1.0

-0.5

0.0

0.5

1.0

1.5

2.0

2.5

3.0

0.00E+00 1.25E-04 2.50E-04 3.75E-04 5.00E-04 6.25E-04

Pote

ntia

l (V)

vs

Ag/

AgC

l



(a) (b)

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

(a) (b)

(c) (d)

Figure 4.2: Corrosion pits observed on the surface of hafnium after cyclicpolarisation 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 RESISTANCEOF 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 propaga-tion curve was also limited by the presence of organic species, which suggestedthat proteins could act as a diffusion barrier limiting the corrosion of hafniumwhen pits have been formed.A drastic reduction of the Eb was observed when imperfections were foundon the material surface (Figure 4.1(b)), indicating a high correlation betweensurface finish and the tendency to suffer from pitting corrosion. Pits were pref-erentially 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 thecyclic polarisation scans. Both CP titanium and hafnium showed a similar be-haviour, that can be divided in four zones with reference to the corrosion poten-tial Ecorr (potential where the net current flow is zero): at potentials below thecorrosion potential (Ecorr), the current density was dominated by the cathodicreaction; at potentials around Ecorr there was a transition from cathodic to an-odic currents; when the potential was increased above the Ecorr, both materialsshowed a passive plateau; finally, when the potentials were further increased,the Eb was reached and the current rapidly increased. In the case of CP tita-nium, the last phase was not observed and the passive layer remained stableat high potentials. Hafnium showed slightly lower values of current density inthe passive state before the breakdown potential was reached.


(a) (b)

Figure 4.4: Evolution of the OCP during the tribocorrosion tests of (a) hafniumand (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 relativelystable OCP value was reached due to the stabilisation of the passivelayer.

• When the tribological test was started at B, the OCP immediately droppedtowards more active OCP values, thereafter a more stable value wasreached indicating the equilibrium between the mechanical depassiva-tion and the electrochemical repassivation.

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

CP titanium showed more noble values than hafnium under the studied condi-tions. Both materials showed the ability to repassivate and rebuild the oxidelayer once the mechanical damage stopped. The grade of repassivation wasevaluated 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 aslightly faster repassivation than CP titanium. The volumetric loss measuredwas higher in the case of hafnium in 0.9% NaCl and 25% serum under thestudied conditions (Figure 4.6), which was influenced by the higher elasticmodulus 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 me-chanical damage stopped in (a) 0.9% NaCl, (b) 25% serum and (c) 50% serumsolution.


0

0.001

0.002

0.003

0.004

0.005

0.006

0.007

0.9% NaCl 25% serum 50% serum

Volumetric

Loss

(cm

3)

Hafnium

Titanium

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

-0.3

-0.25

-0.2

-0.15

-0.1

-0.05

0

0.05

0.1

0 0.5 1 1.5 2 2.5 3 3.5 4 4.5 5

mm

mm

Hf - 0,9% NaClHf - 25% serumHf - 50% serum

-0.3

-0.25

-0.2

-0.15

-0.1

-0.05

0

0.05

0.1

0 0.5 1 1.5 2 2.5 3 3.5 4 4.5 5

mm

mm

CP Ti - 0,9% NaClCP Ti - 25% serumCP Ti - 50% serum

(a) (b)

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

MPa in the case of hafnium and 750 MPa in the case of CP titanium. In addi-tion, the cross section of the wear scar was also deeper in the case of hafniumunder 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 inTable 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 ± 3CP Ti 163 ± 6±: Standard deviation

-1.00

-0.75

-0.50

-0.25

0.00

0.25

0.50

0.75

1.00

-0.0002 0 0.0002 0.0004 0.0006 0.0008 0.001 0.0012

Pote

ntia

l (V)

vs

Ag/

AgC

l



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


Cyclic polarisation tests were performed on hafnium to investigate the effectof fretting on the susceptibility of the material to localised corrosion. Thebreakdown potential of hafnium was dramatically decreased compared to thebreakdown potential in static conditions (Figure 4.8). The effect of the organicspecies contained in the serum solution was similar to that previously reported.Due to the competition between negatively charged proteins and chloride ionsto adsorb on the anodically polarised surface, the breakdown potential washigher 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 wasstudied by monitoring the OCP during the fretting test. The test can be dividedinto three stages as shown in Figure 4.9.

• From A to B the specimen was immersed in the solution and an anodicshift was observed, which was related to the stabilisation of the samplein the solution and thickening of the passive film.


(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 circuitpotential 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 ± 19.4CP 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 valuesand reached a minimum value. After that the OCP recovered and sta-bilised around less active value, which suggested a balance between me-chanical depassivation and electrochemical repassivation.

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

In spite of the more noble OCP values of CP titanium in static condition, thedrop in the depassivation peak (ΔEmax) was more pronounced in the case ofCP titanium in both 0.9% NaCl and 25% serum solution, as shown in Table4.2. Furthermore, the average drop (ΔEavg) in the OCP caused by fretting wasalso higher in the case of CP titanium, suggesting that the passive layer formedon titanium suffered more severe mechanical damage than the oxide layer ofhafnium.

4.3. FRETTING CORROSION OF HAFNIUM 33


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 threezones:

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

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

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

The maximum anodic current (Ipeak), corresponding to the peak observed withthe onset of fretting, and the average anodic current (Iavg) under fretting aresummarised in Table 4.3. A higher peak was observed in the case of CP ti-tanium in both 0.9% NaCl and 25% serum solution, indicating a more severeinitial depassivation caused by fretting. The average current observed for CPtitanium was also higher than for hafnium in the studied conditions. In thecase of hafnium in 0.9% NaCl, the anodic currents reached a relatively stablevalue 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 leadto greater mechanical damage and, subsequently, a more severe depassivation.Furthermore, proteins could have a negative effect on the repassivation of themetals restricting the diffusion of oxygen to the surface [30], which could af-fect not only the corrosion of the material but also the wear induced corrosion,leading to increased wear-corrosion damage.


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 vsAg/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 ± 0.22CP 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 Vin (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 makeit an interesting material for biomedical applications. This work has investi-gated the corrosion behaviour of hafnium in simulated body fluids and theeffect of wear and corrosion on the material degradation:

• Hafnium showed a passive state in simulated body fluids. The oxide filmformed on its surface provided high protection to corrosion.

• Hafnium showed a higher susceptibility to suffer from localised corro-sion than CP titanium. The organic species retarded the formation ofpits and limited their propagation.

• Surface imperfections increased the susceptibility of the material to pit-ting corrosion. Fretting also reduced its resistance to localised corrosion.

• The passive layer of hafnium was disrupted by mechanical loads, lead-ing to higher corrosion. However, the oxide layer was rapidly reformedwhen the damaging action stopped.

• Hafnium showed lower anodic currents than CP titanium when exposedto fretting in simulated body fluids.

Hafnium has shown that the highly protective oxide film formed on its surfacein simulated body fluids provides it with excellent corrosion resistance, evenwhen the material is subject to mechanical damage. However, its main draw-back is the susceptibility of the material to localised corrosion, which is largelyincreased on scratched surfaces and under fretting.

35

Chapter 6

Future Work

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

This work has provided insight into the corrosion and wear corrosion be-haviour of hafnium in simulated body fluids. However, if hafnium is to beimplanted into the body for a particular application, more realistic studies withspecific geometry and loads, similar to those faced in vivo, should be per-formed.

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

37

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42 REFERENCES

Part II

Appended Papers

43

Paper A

Tribology, Corrosion andTribocorrosion of Metal onMetal implants

45

A.1. ABSTRACT 47

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

Tribology, Corrosion and Tribocorrosion of Metal onMetal 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 onpolyethylene implants, especially in case of young patients who require a safeand long-term performance of the device. The reduction of wear particles is akey 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 theartificial joint replacement. However, corrosion of the metallic implant leadsto an increase of the ion levels in the body of the patient. Metallic ions mayproduce a host response that can induce a catastrophic failure of the implant.This review initially focuses on the consequences that the degradation of themetals used in orthopaedic implants have for the health of the patient, and thedifferent 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 aswell. Special ttention is given to the tribology, corrosion and tribocorrosionbehaviour of metal on metal implants. Finally, an overview of mathematicalmodels that have been used to model the behaviour of the implants are alsopresented.

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 surgi-cal interventions of the 20th century due to their effective treatment of chronicpain and diseases such as osteoarthritis, rheumatoid arthritis, hip fractures andtumours. During surgery, the unhealthy hip joint is replaced with an implantin order to restore the function and activity of the articulation. Today, it isestimated 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 spiteof the limitation in manufacturing techniques, some of the first generation im-plants were clinically successful for 20 years or more [2]. However, the in-cidence of failure due to loosening was high. Metal on metal models such asMcKee-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 lev-els, but, due to the high level of success and improvement in surgical methods,the indications for joint arthroplasty have been expanded to younger and moreactive patients [3]. The longevity of MoP implants in the case of young pa-tients is decreased due to the increased activity of the patient. Polyethylenehigh wear rates provoke a macrophage reaction that results in osteolysis andthe loosening of the implant. Revision of the implant is inevitable after a periodof time, which is normally more expensive and complicated than the primaryarthroplasty.Over the last two decades, interest interest has increased in the so called hardon hard bearings, which have been developed to improve the longevity of theimplants and decrease the risk of revision. They present low wear rates, whichcould decrease the risk of a biological reaction that may lead to osteolysis.Ceramic on ceramic articulations are a promising alternative due to their goodbiocompatibility and wear resistance. Despite improvements in the mechan-ical properties of ceramics, structural fracture of ceramic implants remains amajor disadvantage for this type of bearing material [4]. Furthermore, cases ofaudible squeak have been reported [5].The use of different bearing materials (ceramic on metal) has shown reducedwear 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 improvedmaterials and design. Optimised manufacturing techniques allowed improvedclearances 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 andnanoscale wear debris limit the use of this type of implants [9]. The topic hasbeen 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 excellentwear properties, the potential for improved stability and high range of motionwith larger diameter femoral heads [11]. Furthermore, midterm results withcurrent generation implants are promising. Metal on metal implants generateapproximately 100-fold less volumetric wear than metal on polyethylene im-plants [12], the particle size has been estimated to be in the range of 60 to 90nm, 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 ofthe corrosion of the metallic surfaces but also by the dissolution of nano-sizedwear debris. Nano-sized particles are highly reactive and can be corroded eas-ily due to their high surface area to volume ratio [15].Normally, metal on metal implants are made of cobalt-chrome-molybdenumalloys. Cobalt, chromium and molybdenum can be found in the body butthey become toxic depending on the concentration [16, 17]. Safe in vivo con-centration limits still undefined, nevertheless there are some reference valuesestablished by international groups [18, 19]. However, it is thought that theions-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 differentmethods 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 tothe ease of specimen acquisition.The daily output of cobalt in the urine can be used as a surrogate marker of theimplant wear because cobalt is the most abundant metal in the alloy, and it ismore 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 toyield more variable metal concentrations [22].To date, one of the most extended methods to measure and compare metal ion


levels is ion concentration in serum. Strong correlation has been found be-tween measured serum and joint fluid concentrations of metal ions, and alsoa significant correlation between the maximum wear-scar depth and the metalion concentrations in both serum and joint fluid [23]. These correlations leadto the conclusion that serum ion levels of chromium and cobalt can be used toestimate the amount of wear taking place in metal on metal implants.A standard analysis method would be an interesting tool to compare the per-formance of different types of prostheses. The comparative information wouldbe very valuable as long as parameters such as the patient demographics, ac-tivity 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 comingyears.

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

Chromium, cobalt and molybdenum ion levels in patients with metal on metalimplants have been largely compared with other bearing couplings and withpersons without any implant. The analysis of metal ion levels of stable ceramicon ceramic implants shows a negligible ion release in this type of prosthesis[24]. Patients who receive metal on metal implants show significanly highermetal ion levels than patients with metal on polyethylene and non-implantedpersons [25, 26]. In the case of ceramic on metal prostheses the difference islower, 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 arthro-plasty 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 corro-sion is considered related to the area of the bearing surfaces [27] but a largerfemoral 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 greaterexposure of cobalt and chromium ions than small- diameter bearings [29],other recent studies show that there is no difference in metal ion levels betweenpatients with small- or large- diameter femoral heads [30]. This uncertainty isaffecting the introduction of large-diameter metal on metal couples, as used inhip resurfacing.

A.4. PROBLEMS RELATED TO METAL ON METAL IMPLANTS 51

Conservative metal on metal hip resurfacing arthroplasty has been demon-strated as an alternative in young and active patients. The retention of theupper femoral bone stock allows more options at a later time if future revisionof the implant is needed. Furthermore, stress shielding in the proximal femoralshaft is avoided and the risk of dislocation is lower. It has been reported that theincrease in Cr and Co concentrations in the case of large-diameter hip resurfac-ing arthroplasty is significantly higher than the increase in a standard-diameter(28 mm) metal on metal total hip arthroplasty [31, 29] but other studies reportno differences between them [27, 32, 33]. Moroni et al. [34] studied thesedifferences at midterm and they reported no differences in ion levels betweenpatients with either metal on metal Birmingham hip resurfacing or metal onmetal total hip arthroplasty.The effect of the correct implantation must not be overlooked since it has agreat impact on the implant’s performance, especially in the case of hard onhard prostheses. For instance, the abduction angle of the acetabular compo-nent affects the metal ion levels in hip resurfacing components [35]. A moresteeply-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 rangerequired for phagocytosis by macrophages, which eventually leads to a pro-cess of aseptic loosening of the implant [36]. By contrast, particles generatedby metal on metal implants are within the nano-metre range, which reducesmacrophage reaction. However, these nanoparticles cause some other reac-tions 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. Invitro studies show that the biological response of human cells to metal parti-cles depends on the particle size. Nanoparticles are more easily ingested bythe cell than micro-sized particles and they are rapidly corroded, this leads tohigh concentrations of metal within the cell. This results in greater cell deaththan in the case of microparticles [37]. Cobalt nanoparticles present higherin-vitro cytotoxicity than chromium nanoparticles [38]. At the same time, thecytotoxic dose of cobalt nanoparticles is higher than the correlative cytotoxicdose of cobalt ions, but it is still not clear if the combination of the two results


the cytotoxic effects. The concentration range found in vivo in patients withpseudotumours is at least three orders of magnitude lower than the critical con-centration found in vitro [39]. This finding suggests that prolonged exposureat lower concentrations may also affect the cells, which has also been observedin-vitro [1].

A.4.2 Hypersensitivity

Hypersensitivity is described as a cell-mediated lymphocyte-dominated im-munological response [40, 41]. It involves mainly T-cell activation but alsoB-cell activation [42]. Some of the associated symptoms are idiopathic painand aseptic loosening [43]. This has been clinically reported in some arti-cles associated with both hip resurfacing and total hip replacement metal onmetal implants [44, 45, 46, 47]. The frequency of this delayed cell-mediatedresponse is still unknown but appears to be low, it has been estimated to bebetween 0.3% and 5% [48, 49] but it requires more in depth study becauseit leads to the failure of the implant. When this immunological reaction isdetected, the implant must be replaced by another with different bearing mate-rials [40]. The risk of surgical revision does not increase in patients with metalallergies [50], however, metal sensitivity rate in patients showing a peripros-thetic T-lymphocytic inflammation has been reported to be higher than in thegeneral population [41, 51]. Methods to detect potential risk of immunologi-cal response have to be developed, but a postoperative metal sensitivity couldindicate an immunological response to the implant. This reaction seems to beexclusive for metal on metal implants [52], however, it has been suggested thatsimilar 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 as-sociated with the implant which is neither malignant nor infective in nature.Various symptoms have been have been associated with the presence of pseu-dotumours, such as pain, palpable lump, nerve palsy and dislocation. The es-timated incidence is approximately 1% five years post-operation, but it couldincrease 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 causedby the excessive presence of wear particles due to some problem with the ar-ticulation but it could also be due to an allergic reaction to the presence ofmetallic particles. Reported cases occur mainly in the presence of metal onmetal hip resurfacing [53, 56, 57, 58], it could be linked with the higher ionlevels reported by some authors with this type of implants.

A.4.4 Osteolysis

Particulate wear debris produced by metal on polyethylene is responsible forthe macrophage response that causes osteolysis and eventually long-term fail-ure. It is due to the size range of the particles generated. The risk of osteolysisis lower in metal on metal implants because the nano-sized particles generatedgenerally do not produce this biological reaction. Clinical studies on metal onmetal implants have shown mostly good results but there are some cases ofosteolysis [59, 60, 61]. The presence of osteolysis in metal on metal implantedpatients has been associated with hypersensitivity responses to the particlesand ions released by the prosthesis [62, 63, 64]. However it is not fully clearif the hypersensitivity leads to instability of the implant or, by contrary, theinstability of the implant causes excessive wear debris which then leads to themacrophage 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 possibil-ity of genetic effects. Chromium and cobalt present in metal on metal implantsare potentially mutagenic metals [66, 67]. It has been shown that particlesof cobalt chrome alloy can cause chromosome and DNA damage in humancells [68]. The damage caused by cobalt chrome alloy particles is significantlyhigher than other material particles such as alumina [1]. DNA damage wasfound when synovial fluid collected at revision surgery was cultured with hu-man fibroblast [69]. Chromosome aberrations such as chromosomal translo-cations and aneuploidy increase in the peripheral blood of patients within twoyears after metal on metal joint arhroplasty [70]. However, similar increaseis found in the peripheral blood of patients at revision surgery of metal onpolyethylene implants [71] The long-term effects of these chromosome aber-rations are still unknown, further studies are necessary to evaluate the risk andconsequences of implanted metal on metal prostheses.


A.4.6 Cancerogenesis

There are concerns that chromosomal aberrations and DNA damage caused bymetal particles and ions may lead to carcinogenesis, but it is still not clear. Onone 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 isnot higher than the expected in the general population. Which is supportedby more recent studies [73]. On the other hand, the risk of cancer in patientswho undergo metal on metal implants was found to be higher than in patientswith metal on polyethylene implants [74]. Kidney and prostate cancers, lym-phophoma or leukaemia have a statistically higher occurrence in patients withMoM implants [75, 76]. However, the risk of cancer could be higher for pa-tients with rheumatoid arthritis or osteoarthritis [77] due to the inflammatorynature of these diseases. Understanding the long term systemic effects causedby the metal particles and ions released by metal on metal implants is veryimportant for their future use.

A.5 Tribology of metal on metal implants

The study of lubrication, friction and wear in both natural and artificial jointsin biological systems is known as biotribology. In natural joints such as the hipor knee, the bones are covered by cartilage and the articulation is lubricated bysynovial 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. Inspite of the different lubrication modes present under more severe conditions,the deformation of articular surfaces and viscous effects of synovial fluid en-hance the prevalence of elastohydrodynamic lubrication regime, preserving thelow friction and low wear during walking [78].Different types of diseases can affect the properties of synovial fluid as wellas cartilage degradation. Pain can be produced due to direct contact betweenbone surfaces. The replacement of the damaged joint may be necessary insome cases. To date, cartilage cannot be artificially replaced, thus the tribo-logical conditions in the joint vary greatly when a prosthesis is implanted. Thefluid 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

Femoral head

Cartilage

Acetabulum

Ligaments

Synovialmembrane

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 betweenthe asperities of the bearing surfaces and, consequently the friction (FigureA.2) and wear produced on the articular surface. Typically, three differentforms of lubrication are identified.In the boundary lubrication regime the loadis carried by the contact between the asperities of the surfaces. In mixed lu-brication, the order of magnitude of the film thickness is similar to the surfaceroughness, therefore the load is carried by both the asperities and the pressuresgenerated within the lubricant. In the case of full film lubrication, the filmthickness is larger than the surface roughness, there is no asperity contact soideally 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 dominatedfor smaller diameters (16 and 22.225 mm). In this case, reduction in headdiameter results in a proportional reduction of wear. An increase in head sizeof few millimetres, to 28 mm head size, led to a substantial reduction in wearsince the lubrication regime shifts to mixed lubrication. The larger femoralheads studied (36 mm) showed surface separation in some gait phases, whichindicates fluid film lubrication. The corresponding wear rates were very low,


Velocity x ViscosityPressure

Fric

tion

Coe

ffici

ent,

μ

Boundary Mixed Hydrodynamic

Figure A.2: Stribeck curve.

which confirms the beneficial tribological behaviour of larger femoral heads.

Vol

umet

ric w

ear

Increasing diameter of femoral head

boundary

mixed

approaching fluid film

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

A.5.2 Wear phases

The wear of MoM implants presents two different phases, an initial running-inphase characterised by high wear followed by the steady-state phase where thewear rate decreases and reaches a more stable level (Figure A.4). This biphasicbehaviour 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 BIOMEDICALMETALS 57

Number of cycles

Vol

umet

ric w

ear

Running-in Steady-state

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 sur-face, but it can be found in other parts such as the interface neck-head in caseof modular implants. Typically, wear can be classified in three principal pro-cesses: abrasive wear, which can be due to either two or three body, adhesivewear and fatigue wear. Nevertheless other types of wear such as corrosive wearcan be found, especially in presence of metallic surfaces. In vitro studies of thewear mechanisms present in metal on metal artificial joints have shown signsof abrasion, surface fatigue and tribochemical reactions but not adhesion [86].These results have been validated through analysing the surface of retrievedimplants [87].

A.6 Corrosion of biomedical metals

Corrosion plays a key role in the release of metal ions in biological environ-ments. The metallic materials most commonly used for biomedical applica-tions are stainless steel, cobalt chromium alloys and titanium alloys. The cor-rosion resistance of these materials relies on the passive layer formed on theirsurface in contact with corrosive environment.


A.6.1 Passive film composition

The oxidation reaction of passive metals results in the formation of an oxidefilm according to the reaction:

M+nH2O→MOn+2nH++2ne (A.1)

The oxide layer formed on stainless steel and CoCr alloys is primarily com-posed by chromium oxide (Cr2O3). Titanium and titanium alloy layers aremainly composed of titanium oxide (TiO2) [88].

A.6.2 Passive film breakdown

The resistance of the passive layer is important in terms of structural integrityand as well as for the release of degradation products. The susceptibility of thelayer to breakdown depends on the difference between the resting and break-down potential, the more similar these values are the more probable breakdownon small areas of the surface (pitting corrosion) will be. Stainless steels usedfor biomedical applications are likely to suffer from pitting corrosion becausethese two potentials present similar values. In the case of CoCr alloys the dif-ference is higher but corrosion is still possible under some conditions. How-ever, localised corrosion is not common in CoCr alloys, but they typically failby transpassive dissolution [89]. The change in the composition of the passivefilm in the transpassive region leads to an increased dissolution rate, which ismainly dominated by the dissolution of Co. The difference between titaniumand titanium alloys’ potentials is very high so breakdown of the passive layeris not possible under biological conditions [90]. When load and motion arepresent on the surface, the passive layer may be mechanically removed andfretting- 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 thatcan activate the immune system. More active ions will react more easilywith hydroxyl radicals and anions forming oxides and salts, decreasing thepossibility of reacting with biomolecules. Less active ions will stay in thesystem longer and therefore they will have more opportunities to react withbiomolecules and reveal toxicity [88].

A.6.3 Galvanic corrosion

Galvanic corrosion can be caused by the contact between different metals butalso the contact between the same metal with parts of the surface under corro-


sion and parts under tribocorrosion conditions. The contact between differentmetals results in a corrosive attack on the active metal and in relative protec-tion of the more noble metal, which limits the application of some coatings onmetallic materials. This type of galvanic contact is usually present in modu-lar implants, for instance, in the neck-head contact. Under static conditions, agalvanic couple of cobalt alloy and titanium is stable and there is no acceler-ated corrosive attack. However, when stainless steel is combined with eithercobalt alloy or titanium, the couple is unstable and more susceptible to corro-sive attack [91]. In the case of a mismatch between the modular connectionneck-head, relative motion may result in wear on the surfaces, which may af-fect 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 biomaterialsare implanted, therefore the interaction between these proteins and implantedmaterials is an important aspect that should not be neglected. The effect ofproteins on the corrosion depends on material due to the different affinities ofproteins for different materials. Clark et al. [92] studied the effect of the pro-teins serum albumin and fibrogen on some pure materials. In the case of metalssuch as Co and Cu, an increased corrosion rate was reported. The effect on Crand 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 thecorrosion of CoCrMo alloy. The adsorption of albumin on the surface has adouble effect, first as an inhibitor, impeding the access of the oxidant to themetallic surface. On the other hand, most of the proteins present a negativecharge 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 po-tentials albumin accelerates corrosion but around corrosion potential the twoeffects compensate each other [93]. In the case of phosphate ions, its adsorp-tion decreases the anodic reaction rate. Phosphate ions and albumin competeto 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 titaniumis similar to the effect shown on CoCrMo before, the proteins can acceleratethe dissolution of the metal but at the same time they adsorb onto the surfaceblocking the cathodic current. The affinity to adsorption of fibrogen onto ti-tanim surfaces is more than that of albumin [95], but the albumin layer seemsto be more efficient at blocking the cathodic reaction [96]. In the case of tita-


nium alloys, proteins can enhance corrosion resistance of Ti6Al4V, but reducethe corrosion resistance of other alloys such as Ti13Nb13Zr and Ti6Al7Nb[97].The dissolution of metals in presence of proteins is also reported for stain-less steel. The dissolution of the nickel and molybdenum is increased bythe presence of fibrogen. Transferrin and albumin increase the dissolution ofchromium [98].

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

Equation A.2 has been used to describe the total material loss produced intribocorrosion systems [99]

T =W +C′+S (A.2)

where W is the material loss due to wear in absence of corrosion, C’ is thematerial loss due to corrosion in absence of wear, and S represents the synergywhich includes the effect of wear on corrosion and the effect of corrosion onwear.

A.6.5 Tribocorrosion measurements

Tribological tests are commonly combined with electrochemical experimentsin order to study the tribocorrosion. Three-electrode cells are normally in-tegrated in different types of tribometers to monitor the electrochemical be-haviour under tribological contacts. A potentiostat controls and monitors the


Metal nanoparticles Metallic ions

Passivated metal

Mechanical Depassivation

Electrochemical Repassivation

M+ M+

M+ M+

M+

Figure A.6: Schematic representation of the depassivation/repassivation pro-cess.

electrochemical signals of the electrodes while the load and speed are con-trolled by the tribometer.

Tribometer

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

Electrochemical cell

Electrochemical techniques are used to understand the effect of corrosion in thecomplex system. The open circuit potential gives limited information on thekinetics of surface reactions [107]. However, information about the corrosionstate of the material (active or passive) can be obtained, as well as informationabout how the tribological contact affects the passivity of the material and howthe passive layer is recovered once wear ceases. The potential shift can beused as an indicator of the damage of the passive layer (Figure A.7). AnodicPolarisation provides information about the rate of dissolution in a given envi-ronment [99]. The breakdown potential indicates the resistance of the passivelayer to localized corrosion initiation. However, the non-stationary conditionsof the test limit the physical interpretation and quantification of the results. Inpotentiostatic tests the current is measured at fixed potential as a function oftime. The mechanical depassivation and the electrochemical repassivation canbe evaluated following the charge transfer rate. Cathodic protection tests are


Pot

entia

l (V

)

Active material

Loading Unloading

Time (s)

Passive material

Figure A.7: Typical open circuit potential for passive material and active ma-terial 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 andwithout 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 withthe proteins present in biological fluids, which will affect the performance ofthe 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, initiallymore abundant proteins dominate the process but they are replaced over timewith proteins which are more affine for the implant surface [109]. While theexact bonding mechanisms are not fully understood, the presence of high num-ber of free ions close to the surface could ease the formation of metal/proteincomplexes [110]. The stable layer acts as a solid lubricant between the bear-ing surfaces which produces beneficial effects for the wear behaviour, for in-stance the prevention of adhesive wear by avoiding direct metal-metal contact[86, 110]. The tribofilm acts as an effective charge transport barrier that limitsthe ion release [101, 106]. On the other hand, an increase in the dissolutionrate [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 studiedit 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 com-position. It was described as an organometallic film which contained Co andCr. The thickness of the film was about 15-80 nm and it was similar in bothsamples. Similar results have been reported using different test configurations[87]. Recently, signs of a graphitic layer have been found on metal on metalbearing surfaces. However, the mechanisms of formation of such material arenot clear [111]. The formation of the so called tribofilm should be further stud-ied, the mechanical properties should be determined and how it influences thetribology 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 hipimplants has a great influence on the amount of degradation products releasedinto the body. If full film lubrication is achieved, the direct contact betweensurfaces is minimized, which potentially reduces wear debris. The lubricationof MoM implants has been widely investigated using computational simulatingtechniques, that are a powerful tool for understanding problems and predictingresults. Mathematical models have been developed to study the tribologicalperformance of the bearing in MoM prostheses.

Elastohydrodynamic lubricaion

NOTATIONcd radial clearance Rc cup radiuse eccentricity Rh head radiusE ′ Young modulus t timef calculated load u entraining velocityh film thickness w applied loadhmin minimum film thickness δ elastic deformationp pressure η viscosityR equivalent radius = 1/Rh−1/Rc

Due to the geometry of the femoral head-cup contact, spherical coordinatesare normally used to establish the main equations governing the EHL problems(Figure A.8) [112].


z

y

x

P

Figure A.8: Spherical coordinates.

The pressure of the fluid film in the bearing is governed by the Reynoldsequation (Equation A.3).

∂∂ϕ

(h3∂p∂ϕ

)+ sinθ∂

∂θ(h3 sinθ

∂p∂θ

) =

6ηR2c sinθ

⎡⎢⎣−wx(sinϕsinθ∂h

∂θ + cosϕcosθ ∂h∂ϕ

+wy(cosϕsinθ∂h∂θ − sinϕcosθ ∂h

∂ϕ+wz sinθ ∂h

∂ϕ

⎤⎥⎦

+12ηR2c sin2 θ∂h∂t

(A.3)

The film thickness is usually expressed as shown in Equation A.4, where bothrigid and elastic deformation between the two bearing surfaces are considered[112].

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

The elastic deformation of the surfaces is modelled by Equation A.5, where Kdenotes the influence coefficient of the elastic surfaces and θm denotes a fixedmean latitude [112].

δ(ϕ,θ) =∫

ϕ

∫θK(ϕ−ϕ′,θ−θ′,θm)p(ϕ′,θ′)dθdϕ (A.5)

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).

fx,y,z = R2c∫

ϕ

∫θpx,y,zdθdϕ = wx,y,z

⎧⎨⎩px = psin2 θcosϕpy = psin2θsinϕpz = psinθcosθ

⎫⎬⎭

(A.6)

The two different models primarily used to perform elastohydrodynamic anal-yses of metal on metal hip joint replacements are the ball-in-socket and the ballon plate (Figure A.9). Although the geometric similarity with a real hip jointis higher in the ball-in-socket lubrication model, it has been demonstrated thecapability of both to perform steady-state EHL analysis [113, 114]. The maindifference between them was found for large cup inclination angles, where theball-in-socket model predicted no change in film thickness while the ball onplane model predicted reduction in film thickness due to the lubricant supplyat the cup edge. The approximation of synovial fluid with Newtonian, iso-viscous and incompressible fluid is normally used because it does not showsignificantly different results under typical operating conditions in comparisonwith more sophisticated non-Newtonian model that considers the shear thin-ning effect of synovial fluid [113]. Simulated operating conditions influencethe results of the lubrication model. Elastohydrodynamic lubrication analysishas been performed under transient conditions, such as a walking cycle or thecycling gait [115, 116, 117]. Results obtained simulating the whole walkingcycle show different film thickness in comparison with steady-state conditions[112], which shows that the results are highly influenced by the loading andmotion conditions simulated and highlights the importance of realistic param-eters.


Mixed lubrication

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

A.7.2 Wear prediction

Computational methods have been also used to simulate the long-term effecton wear for metal on metal hip implants and to understand the parameters thatmay affect them during their life [119]. Wear models are normally reproducedas 3D ball-in-socket geometry. The model is normally governed by Archard’swear law (Equation A.7), which calculates the volumetric wear (V) dependingon the resultant normal load (W), the relative surface sliding distance (L), anda coefficient of wear (Kw).

V = KwWL (A.7)

Despite the incomplete description of wear mechanisms of Archard’s wearlaw, it has been used due to its simplicity and validity in wide applications.The models are dependent on the wear coefficient, which is usually obtainedexperimentally from wear testing using a pin on plate or hip simulator to adaptthe model to the characteristics of the studied bearing. The variation of thewear coefficient between the running-in and steady state phase must be care-fully considered.The relationship between metal on metal volumetric wear and elastohy-

drodynamic film thickness was studied by Dowson [119], based on a numberof hip simulator studies performed in different laboratories. The Hamrock-Dowson equation (Equation A.8) was used to derive the Equation A.9, theminimum film thickness depending on the diametral clearance and the head

A.7. MODELLING 67

diameter.Hmin =

hminR

= 2.8( ηuE ′R

)0.65( wE ′R2

)−0.21(A.8)

hmin = 1.6904 ·108 d2.19

c0.77d(A.9)

The experimental wear rates obtained from a number of simulator studies areplotted against the theoretical film thickness in Figure A.10. It is shown that thesteady state wear rate is almost independent of the film thickness whereas therunning in depends greatly on the film thickness. Low running-in and steadystate 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 manyyears 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 tomodel the parameters that control the problem. A number of different ap-proaches have been used by different authors. Mischler et al. [120] proposeda model that describes the anodic corrosion current (Ir) of passive metals slid-ing against a rigid non-conductive body as function of experimental parame-ters such as frequency ( f ), wear track length (L), load (W ) and hardness (H),as well as a proportionality coefficient Kr that cannot be theoretically deter-mined(Eq. A.10).

Ir = Kr L f W 0.5H−0.51/ f∫

0

i dt (A.10)

This model considers the contact between asperities, which is described as afunction of load and hardness, to determine the depassivated area. The predic-tion of the effect of mechanical and physical parameters shows a good corre-lation correlation with the experimental results, however, the wear componentwas not modelled in this work. A model based on the concepts of micro-cracking and fatigue to model the wear related damage, that considers bothcorrosion, wear and the synergism between them was presented by Jiang andco-authors [121]. However, although several tribocorrosion models have beenpublished in the literature, there is a need of bio-tribocorrosion models, wherethe kinetics and mechanisms that affect the process are influenced by the bio-logical environment.


(a)

Elasto-hydrodynamic film thickness (nm)

Run

ning

-in w

ear (

mm

3 )

(b)

Figure A.10: Steady state (a) and running-in (b) wear rates depending on thetheoretical 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 ofthe particles released into the body, which ideally avoids the host responseproduced in the case of metal on polyethylene implants and, consequently,osteolysis and loosening of the implant. The longevity of the implant can betheoretically improved. However, the advantages of metal on metal implantsmay turn into disadvantages. The levels of metallic ions are increased in thebody and these ions do not remain in the area close to the implant but they aredispersed around the body.

• While in some patients the high levels of metallic ions do not have anyconsequence, in some other patients it can lead to different biologicalreactions that can cause pain and failure of the implant. The causesof these types of host reactions are not fully understood. Prediction ofwhich patients are more likely to suffer from severe biological reactionsshould 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 andwear particles. The synergism between them in biological environmentsneeds to be better understood and studied.

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

• Newmaterials with improved tribological and corrosion properties whichlead to decreased degradation products and improved outcomes are de-sirable.

In spite of the concerns related with metal on metal hip implants, theystill are a potential alternative for total joint replacements, therefore, more ef-forts should be made to understand and predict their behaviour and biologicalresponse. Achieving this requires strong collaboration between a variety ofscientific 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 onthe manuscript.

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Paper B

Corrosion and Tribocorrosionof Hafnium in Simulated BodyFluids

79

B.1. ABSTRACT 81

Submitted for journal publication

Corrosion and Tribocorrosion of Hafnium in SimulatedBody Fluids

J. Rituerto Sin1, A. Neville2 and N. Emami1


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, how-ever, little is known about its resistance to wear and corrosion in biological en-vironments. The corrosion and tribocorrosion behaviour of hafnium and com-mercially pure (CP) titanium in simulated body fluids were investigated usingelectrochemical techniques. Cyclic polarisation scans and open circuit poten-tial measurements were performed in 0.9% NaCl solution and 25% bovine calfserum solution in order to assess the effect of organic species on the corrosionbehaviour of the metal. A pin-on-plate configuration tribometer and a threeelectrode electrochemical cell were integrated to investigate the tribocorro-sion performance of the studied materials. The results showed that hafniumhas good corrosion resistance. The corrosion density currents measured in itspassive state were lower than those measured in the case of CP titanium, how-ever, it showed a higher tendency to suffer from localised corrosion, which wasmore acute when imperfections were present on the surface. The electrochem-ical breakdown of the oxide layer was retarded in the presence of proteins.Tribocorrosion tests showed that hafnium has the ability to quickly repassivateafter the oxide layer was damaged; however, it showed higher volumetric lossthan 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 withtitanium and zirconium. The metal was first identified by Dirck Coster andGeorges de Hevesy in Copenhagen in 1923 [1] and owes its name to ’Hafnia’,the Latin name for Copenhagen. Hafnium is always found associated with zir-conium in ores and is primarily found in zircon, with a ratio Hf/Zr of about2.5% [2]. Due to a number of interesting properties such as high ductility andstrength, as well as resistance to corrosion and mechanical damage [3], it hasattracted interest for various applications. For instance, it is used as a controlmaterial for nuclear reactors and as an alloying element in some superalloysused in aircrafts engines [4].In 1984, Marcel Pourbaix included hafnium among the group of metals to beconsidered 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 bodyat that time it was discarded from the final list of 13 metals to be theoreticallyconsidered [5]. The passive layer found on hafnium’s surface, mainly formedof HfO2, is highly stable in neutral solutions [6].More recently, the properties of hafnium as an implant material have been in-vestigated [7, 8]. Matsuno et al. [7] implanted a number of refractory materi-als (titanium, hafnium, niobium, tantalum and rhenium) in subcutaneous tissueand femoral bone marrow of rats, and performed histological observation andelemental mapping in order to evaluate the biocompatibility and osteogenesisof such materials. The results of the in vivo study showed that hafnium hadgood biocompatibility and osteogenesis. A different study on the tissue re-sponse to hafnium [8] showed similar response in soft and hard tissues in twodifferent animal species (rat and rabbit), which suggests that hafnium could bean 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 den-tal implants, bone plates and hip and knee implants. Titanium-6%Aluminum-4%Vanadium (Ti-6Al-4V) is usually the material of choice. However concernsregarding the use of aluminium -which has been pointed out as a growth in-hibitor of bone and possible cause of Alzheimer’s disease [9]- and vanadium-which can lead to cardiac and renal dysfunction [10]- enhance the need toinvestigate new alloying elements. Different Ti-Hf binary alloys have been re-ported in the literature. Wrought Ti-Hf alloys have shown a gradually increasein yield and tensile strength with increased amount of hafnium, up to 40%, atthe 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-Hfalloys [12]. The electrochemical behaviour of this type of alloys is similar topure titanium, showing excellent corrosion resistance [13].Titanium-Niobium-Hafnium-Zirconium (Ti-Nb-Hf-Zr) alloys have also beeninvestigated [14]. These alloys have shown a low elastic modulus which isbeneficial in order to reduce the stress shielding effect and to enhance bonegrowth. It has been also shown that cold work can be used to decrease theelastic modulus of this type of alloys, reaching values close to the elastic mod-ulus of cortical bone [15].To date, the behaviour of pure hafnium in a biological environment has notbeen studied in great depth. Little is known about the resistance of the passivelayer under wear-corrosion conditions and the effect of proteins on its corro-sion and tribocorrosion behaviour. Thus, there exists a need for further studiesof hafnium under biological conditions in order to determine the suitability ofthis material for biomedical applications. The present work studies the cor-rosion resistance of hafnium in simulated body fluids as well as the effect ofwear 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(GfEMetalle und Materialien GmbH, Germany) were used in this study. Plates6 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 sealedwith silicon to avoid the presence of crevices. The plates were polished withsilicon carbide (SiC) paper up to 1200 grit followed by a 9 µm and a 3 µmdiamond paste. Finish polishing was accomplished with a colloidal silica sus-pension. Specimens were cleaned and rinsed with distilled water and ethanolbefore they were assembled into the measurement cell. Two different surfacefinishes were investigated in the case of hafnium: a first group of specimenswere studied immediately after polishing; a second group of specimens werescratched after polishing, to assess the effect of surface imperfections on thecorrosion resistance of hafnium.In order to simulate body fluids, 25% and 50% bovine calf serum (HarlamSera-LabÂo, UK) solutions were employed. In addition, 0.9% NaCl solutionwas 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 .


B.3.2 Electrochemical tests

All electrochemical measurements were carried out using an Autolab PG-STAT302N potentiostat (Metrohm Autolab, The Netherlands). A three elec-trode electrochemical set-up was used. A platinum wire was used as a counterelectrode, 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. Allpotentials are given with respect to the Ag/AgCl electrode. Specimens wereimmersed 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 currentdensity was limited to a maximum of 500 µA cm-2. All tests were repeated atleast three times to ensure repeatability. Optical microscopy (Olympus Vanox-T, Olympus Optical, Japan) was used to analyse the surface of the specimensbefore and after measurements.

B.3.3 Tribocorrosion tests

A tribocorrosion cell compromising of a three electrode electrochemical celland a reciprocating tribological tester was used. All electrochemical measure-ments were carried out using a potentiostat VersaSTAT 4 (Princeton AppliedResearch, USA). The three-electrode cell configuration is the same as the ear-lier described, with an Ag/AgCl electrode as the reference electrode, a Pt wireas counter electrode and the specimen as the working electrode. The slid-ing wear tests were performed in a reciprocating wear tester that conformsto ASTM G133. The studied materials were rubbed against a silicon nitrideball with 12 mm diameter. A normal load of 40 N was used. The frequencywas controlled at 1 Hz and the stroke length was set at 10 mm. Specimenswere passivated in 30% HNO3 for 30 minutes before each test. The mass ofthe specimens was gravimetrically determined by using a digital microbalance(Mettler Toledo AT201, UK) with 0.01 mg resolution. Each pin was weighedbefore and after each test at least five times to ensure repeatability. Specimenswere immersed in the selected solution in static conditions before the onset ofthe tribological test and the OCP was monitored for 60 minutes. After thatthe tribological test was started and the OCP was measured for 60 minuntes.Finally, the sliding was stopped and the OCP was measured for 60 minutesto study the repassivation kinetics of the materials in different environments.3D optical profilometry (Wyko 1100NT, Veeco Instruments, USA) was used tostudy 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

-1.0

-0.5

0.0

0.5

1.0

1.5

2.0

2.5

3.0

0.0E+00 5.0E-04 1.0E-03 1.5E-03 2.0E-03 2.5E-03 3.0E-03

Pote

ntia

l (V)

vs

Ag/

AgC

l



-1.0

-0.5

0.0

0.5

1.0

1.5

2.0

2.5

3.0

0.00E+00 1.25E-04 2.50E-04 3.75E-04 5.00E-04 6.25E-04

Pote

ntia

l (V)

vs

Ag/

AgC

lCurrent 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 solutionand 25% serum solution are shown in Figure B.1(a). A sudden increase of thecurrent density was observed in both solutions due to the breakdown of thepassive layer. The breakdown potential increased in presence of proteins, asshown in Table B.1. Once the current density reached the current limit (500µA cm-2), the potential slope was reversed, however the current density keptincreasing after the reversal point was reached, showing typical pit propagationbehaviour. The pit propagation curve was larger in 0.9% NaCl solution. Therepassivation potential in the case of 25% serum solution was slightly higherthan in the saline solution. As shown in Figure B.1(b), the breakdown potentialshowed a large decrease when the surface of the material was scratched.Formation of pits due to localised corrosion was observed under optical mi-croscopy. The pits observed on the polished specimens (Figure B.2(a) andB.2(b)) were larger than in the case of the scratched surfaces, where, small pitswere preferentially formed on the surroundings of the surface imperfections(Figure B.2(c) and B.2(d)).


(a) (b)

(c) (d)

Figure B.2: Corrosion pits observed on the surface of hafnium after cyclicpolarisation 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 passiveregion at different applied potentials.

Hafnium CP titanium0.9% NaCl 25% serum 0.9% NaCl 25% serum

Potential (mV) Passive Current Density (µA cm−2)0 4.26 ± 0.18 5.65 ± 0.02 5.95 ± 1.52 11.99 ± 0.24250 5.36 ± 0.03 6.08 ± 0.05 9.66 ± 0.79 11.17 ± 0.16500 6.25 ± 0.25 6.43 ± 0.01 10.21 ± 0.02 10.90 ± 0.31750 6.97 ± 0.38 6.78 ± 0.19 10.31 ± 0.49 10.46 ± 0.501000 7.44 ± 0.38 7.15 ± 0.10 10.2 ± 0.96 10.28 ± 0.071250 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 in0.9% NaCl and 25% serum solution, which could be divided in four zones: atpotentials below the corrosion potential (Ecorr), the current density was domi-nated by the cathodic reaction; at potentials around the Ecorr there was a transi-tion from cathodic to anodic currents; when the potential was increased abovethe Ecorr, both materials showed a passive plateau where a stable oxide layerwas formed on their surface; finally, when the potential was further increasedthe passive film was disrupted and the current rapidly increased. In the case ofCP titanium, breakdown of the oxide layer was not observed, the passive layerremained stable at high potentials. The current densities of hafnium showed


Table B.3: Tafel extrapolation of the corrosion potential and corrosion currentdensity in 0.9% NaCl and 25% serum.

Hafnium CP titanium0.9% NaCl 25% serum 0.9% NaCl 25% serum

Ecorr (V)-0.638 -0.881 -0.415 -0.613± 0.016 ± 0.045 ± 0.011 ± 0.051

icorr (nA cm−2) 163.9 15.3 395.9 146.0± 40.9 ± 3.5 ± 114.9 ± 32.0


-1.6

-1.4

-1.2

-1.0

-0.8

-0.6

-0.4

-0.2

0.00 3600 7200 10800

OC

P (V

) vs

Ag/

AgC

l

Time (s)

Hafnium

0.9% NaCl25% serum50% serum

-1.6

-1.4

-1.2

-1.0

-0.8

-0.6

-0.4

-0.2

0.00 3600 7200 10800

OC

P (V

) vs

Ag/

AgC

l

Time (s)

Titanium

0.9% NaCl25% serum50% serum

(a) (b)

Figure B.4: Evolution of the OCP during tribocorrosion test on (a) hafniumand (b) CP titanium.

slightly lower current density values than titanium in its passive state (TableB.2) before the breakdown potential was reached. The values of corrosion po-tential (Ecorr) and corrosion current density (icorr) of titanium and hafnium in0.9% NaCl solution and 25% serum solution extracted from the Tafel plots fit-ted to the potentiodynamic curves are shown in Table B.3. Hafnium showeda lower Ecorr than titanium in both solutions. The icorr observed for hafniumwas slightly lower than that for CP titanium in both 0.9% NaCl solution and25% 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, whenthe specimens were immersed in the electrolyte solution. With the onset ofthe 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 mechanicaldamaging 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, indicatingthat the oxide layer was rebuilt on the surface.The evolution of the OCP of hafnium and CP titanium after mechanical de-passivation ceased is shown in Figure B.5. The percentage of repassivataionwas calculated according to Equation (1).

Repassivation(%) =Et −Ewear

Epassive−Ewear (B.1)

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


0

0.001

0.002

0.003

0.004

0.005

0.006

0.007

0.9% NaCl 25% serum 50% serum

Volumetric

Loss

(cm

3)

Hafnium

Titanium

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

-0.3

-0.25

-0.2

-0.15

-0.1

-0.05

0

0.05

0.1

0 0.5 1 1.5 2 2.5 3 3.5 4 4.5 5

mm

mm

Hf - 0,9% NaClHf - 25% serumHf - 50% serum

-0.3

-0.25

-0.2

-0.15

-0.1

-0.05

0

0.05

0.1

0 0.5 1 1.5 2 2.5 3 3.5 4 4.5 5

mm

mm

CP Ti - 0,9% NaClCP Ti - 25% serumCP Ti - 50% serum

(a) (b)

Figure B.7: Wear scar cross-section profiles measured after the tribocorrosiontests on (a) hafnium and (b) CP titanium.

before the mechanical damaging starts. After the first quick increase, the po-tential kept rising at a lower rate due to the growth of the protective film. TheOCP increased slightly faster in the case of hafnium in all fluids.The measured volumetric losses of hafnium and titanium in 0.9% NaCl solu-tion, 25% serum and 50% serum are shown in Figure B.6. Hafnium showedhigher volumetric loss in 0.9% NaCl solution and 25% serum solution. Thevolumetric loss was generally higher for 25% serum than for 0.9% NaCl solu-tion, however, it did not increase in the case of 50% serum in comparison with25% serum. The depth of the wear scars measured for hafnium was larger inthe case of hafnium (Figure B.7(a)) than for CP titanium (Figure B.7(b)) in allsolutions.

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 affectedby different factors such as infections or surgery, where the pH can vary fromvalues around 5.6 up to 9 [5]. A preliminary study of the potential-pH curvesof hafnium and titanium showed that these two materials are in passive statein the body’s pH range [5]. Hafnium is spontaneously passivated in aqueousenvironments by forming a hafnium oxide layer on the surface [6]. Titaniumis well known to be in the passive state in biological environments. A titaniumoxide layer provides this material with excellent corrosion resistance. Titaniumand hafnium are generally passivated by the reaction with water according tothe equation (1) and (2) respectively.

Ti+2H2O↔ TiO2+4H++4e− (B.2)

H f +2H2O↔ H fO2+4H++4e− (B.3)

In simulated body fluids, such as the studied solutions, both titanium andhafnium formed a stable passive layer and OCP values were in the passivedomain. The oxide layer formed on hafnium has been shown to be very pro-tective, showing slightly lower current densities than those for titanium. Bothmaterials also showed low corrosion current density in the presence of organicspecies. The icorr values extracted from the Tafel plots showed lower corro-sion currents densities in 25% serum solution, indicating that proteins mightbe acting as a cathodic inhibitor [16], however, no significant difference wasobserved in the corrosion current densities measured in the passive region inthe saline and in the serum solutions.

Cyclic polarisation tests revealed that hafnium tends to suffer from localisedcorrosion at higher potentials. The results showed that the presence of organicspecies increased the breakdown potential of the film, which could be dueto the positive charge of proteins at the pH of the test solutions (7.4) [17].In the polarisation scan, specimens were anodically polarised, therefore thenegatively charged species tended to be attracted towards the surface of thespecimen. Proteins would compete with chlorides to adsorb on the surface,which would retard the formation of pits. Similar effects have been reportedfor AISI 316L stainless steel [18]. The breakdown potential measured in both0.9% serum and 25% serum solution exceeded the anodic oxygen evolutionpotential, however this reaction could not be detected, due to the insulating


effect of HfO2 for anodic reactions [19]. The pit propagation curve was alsolimited in the presence of organic species. Proteins would be attracted towardsthe region were positively charged ions were released [18]. The presence ofproteins could be act as a diffusion barrier; therefore the current density valuesreached when the pits were formed in serum solution were lower than in thecase of saline solution. A strong correlation between the surface finish and re-sistance to localised corrosion was also observed. When surface imperfectionswere present on the surface of hafnium, its resistance to localised corrosionwas greatly decreased, and corrosion pits were formed around the surface im-perfections.

B.5.2 Tribocorrosion of hafnium

As discussed earlier, both hafnium and titanium show a passive state whenimmersed in aqueous solution. The passive state was also observed in thepresence of organic species, which provided them with excellent corrosion re-sistance. The mechanical damage to which the materials were exposed causeddamage to the passive layer and the bulk material to be in contact with themedia, therefore the OCP dropped towards more active values. When the me-chanical damage ceased, both materials had the ability to repassivate, and theoxide layer was quickly reformed on their surface, which could be observedby a quick increase on the OCP. Hafnium showed a slightly faster repassiva-tion that was observed by a faster increase of the OCP in all solutions. Thepresence of organic species slowed down the reformation of the passive layerin both cases.The volumetric loss measured in the case of hafnium was higher in 0.9% NaCland 25% serum under the studied conditions. This was influenced by the higherelastic modulus of hafnium (138MPa [20]) as compared to CP titanium (102.7MPa [21]), which yields an initial Hertzian contact pressure of approximately850 MPa in the case of hafnium and approximately 750 MPa in the case of CPtitanium in this study. However, the volumetric loss decreased in 50% serumfor both materials in comparison with 25% serum solution, which suggestedthat the increase of the protein content of this solution may improve the lubri-cation of the system [22, 23].

B.6 Conclusions

Hafnium has been pointed out as an interesting material for biomedical ap-plications due to its good biocompatibility and osteogenesis [8]. However its

B.7. ACKNOWLEDGEMENTS 93

corrosion and tribocorrosion behaviour in the biological environment have notbeen fully studied. In the present study, the corrosion and tribocorrosion be-haviour of hafnium have been investigated using CP titanium as the referencematerial.

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

• Hafnium had a higher tendency to suffer from localised corrosion thanCP titanium. The presence of proteins retarded the formation of pits andreduced their propagation.

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

• The wear rate shown by hafnium was higher in comparison to CP tita-nium in 0.9% NaCl solution and 25% solution under the studied condi-tions.

B.7 Acknowledgements

J. Rituerto would like to acknowledge the Swedish Research of Tribology forsupporting the visit to the University of Leeds, UK. This work was fundedby 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.


[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




[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. Okunoand 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.



[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 volume13B: 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 ofMaterials.

Paper C

Fretting Corrosion of Hafniumin Simulated Body Fluids

97

C.1. ABSTRACT 99

To be submitted

Fretting corrosion of Hafnium in Simulated Body Fluids

J. Rituerto Sin1, A. Neville2 and N. Emami1


2Institute of Engineering Thermofluids, Surfaces and Interfaces,School of Mechanical Engineering, University of Leeds, Leeds LS2 9JT UK

C.1 Abstract

Metals used in medical devices can be exposed to fretting in applications suchas dental and joint implants. The effect of fretting, combined with the cor-rosiveness of body fluids, leads to enhanced degradation that can eventuallyresult in failure of the implant. Hafnium has been suggested as an interest-ing material for biomedical applications due to its good biocompatibility andosteogenesis. However, its behaviour in fretting corrosion conditions has notbeen studied in depth. In order to investigate the effect of fretting on the corro-sion resistance of hafnium and commercially pure (CP) titanium in simulatedbody fluids, a three electrode electrochemical cell integrated with a ball onflat reciprocating tribometer was used. The effect of fretting on the resistanceof hafnium to localised corrosion was investigated. The susceptibility to lo-calised corrosion was drastically decreased when hafnium was subjected tofretting. The study of the open circuit potential showed a more severe mechan-ical depassivation due to fretting in the case of CP titanium in comparison tohafnium. In addition, the anodic currents measured during potentiostatic testswere also higher in the case of CP titanium.

Keywords: Hafnium, Titanium, Biomaterial, Corrosion, Fretting corro-sion

100 PAPER C. FRETTING CORROSION OF HAFNIUM

C.2 Introduction

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

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

Hafnium is a transition metal with high ductility and strength, and goodresistance to corrosion and mechanical damage [9]. A very protective oxidelayer, mainly formed by HfO2 [10], provides excellent corrosion resistance.Various studies have shown that hafnium has good biocompatibility and os-teogenesis [11] and a tissue response similar to that observed for titanium[12]. However, there is limited information about the behaviour of hafniumin biological environments. Despite the passive state and the good corrosionbehaviour of the material in simulated body fluids, a tendency to suffer fromlocalised corrosion has been reported in previous studies [13]. To date, the be-haviour of hafnium in biological environments in fretting-corrosion conditionshas not been studied in depth. In the present study, electrochemical techniqueshave 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 inbiomedical 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(GfEMetalle und Materialien GmbH, Germany) were used in this study. Platesof 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 fol-lowed by a 9 µm and a 3 µm diamond paste. Colloidal silica suspension wasemployed for the final polishing. Samples were cleaned and rinsed with dis-tilled 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 toisolate the effect of the organic species (proteins, amino acids, etc.) from thesaline environment. All tests were conducted at 37◦C .

C.3.2 Microhardness measurements

A Matsuzawa MXT-CX microhardness tester (Matsuzawa Co., Japan) wasused to study the microhardness of CP titanium and hafnium. At least fivemeasurements were performed on each sample to ensure repeatability.


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 mountedon a polyoxymethylene holder and were fretted against a stationary silicon ni-tride ball (grade 10) of 12 mm diameter (Redhill Precision, Czech Reublic). Anormal load of 10 N was used. The frequency was controlled at 1 Hz and themaximum stroke length was 120 µm. Tests were performed for 3600 cycles. Athree electrode electrochemical cell was integrated into the system to performthe fretting corrosion tests. The schematic diagram of the fretting corrosion setup is shown in Figure C.1. The electrochemical cell consisted of a platinumwire used as the counter electrode, an Ag/AgCl electrode (3M KCl) used asthe reference electrode, and the studied material was used as the working elec-trode. All potentials are given with respect to the Ag/AgCl electrode, which is196 V on the standard hydrogen electrode scale. Three different electrochem-ical tests were performed in this study:

Cyclic Polarisation Tests

Cyclic polarisation scans were performed on hafnium in fretting conditionsin 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 wasreached. Samples were polarised in the cathodic direction down to -1 V at ascan rate of 10 mV s−1. Finally, the cyclic polarisation test was performed at ascan rate of 2 mV s−1 up to a potential of 3 V or until the current density limit(800 µA cm−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 testconsisted of three different phases. First, the sample was immersed in staticconditions for 3600 s. Then, the specific load was applied and the fretting wasstarted. Finally, after 3600 cycles fretting was stopped and the sample wasunloaded. Tests were repeated at least three times.

Potentiostatic Tests

A constant passive potential of 0 V (vs Ag/AgCl) was applied to monitor thecurrent as a function of time and the effect of the mechanical depassivation dueto 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 ± 3CP Ti 163 ± 6±: Standard deviation

Table C.2: Comparison of the breakdown potential of hafnium in fretting con-ditions (E f rettingb ) and in static conditions (Estaticb ) in 0.9%NaCl and 25% serum.

E f rettingb (V) Estaticb (V) [13]0.9% NaCl 0.279 ± 0.003 1.387 ± 0.08025% serum 0.405 ± 0.007 2.331 ± 0.128


at 0 V during 3600 s before the start of fretting; after that, the sample wasfretted during 3600 s; finally, fretting was stopped and currents were monitoredfor 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. Hafniumshowed 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 ofpits on the surface. The formation of the pits and the subsequent localisedcorrosion were retarded in 25% serum (Table C.2). Once the potential scanwas reversed, the pit propagation caused that the currents continued increas-ing. The currents reached in the saline solution were higher than in the serumsolution. The repassivation potential, at which currents reached the originalpassive 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


-1.00

-0.75

-0.50

-0.25

0.00

0.25

0.50

0.75

1.00

-0.0002 0 0.0002 0.0004 0.0006 0.0008 0.001 0.0012

Pote

ntia

l (V)

vs

Ag/

AgC

l



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

Table C.3: Maximum drop (ΔEmax) and average variation (ΔEavg) of the OCPdue to mechanical depassivation caused by fretting.

ΔEmax (mV) ΔEavg (mV)

Hf 0.9% NaCl 79.2 ± 25.9 36.9 ± 1.625% serum 151.9 ± 36.3 39.9 ± 19.7

CP Ti 0.9% NaCl 227.3 ± 43.6 81.3 ± 5.525% serum 260.4 ± 22.0 171.9 ± 21.3


(a) (b)

Figure C.3: Effect of fretting on the open circuit potential of hafnium and CPtitanium 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 0Vin (a) 0.9% NaCl and (b) 25% serum.

three phases: from A to B (in Figure C.3), before fretting started, an anodicshift was observed which suggested a thickening of the passive layer; whenfretting started at point B (in Figure C.3), the OCP values dropped towardsmore active values. The OCP reached a minimum value immediately after theonset and it tended to increase towards less active values due to the equilib-rium between the mechanical depassivation and electrochemical repassivation.Once the fretting was stopped at point D (in Figure C.3), the OCP recoveredtowards more positive values indicating repassivation of the material. The val-ues of the maximum variation of the OCP and the average variation duringfretting are shown in Table C.3. Despite titanium showed more noble OCPvalues in static conditions, the variation of the values due to fretting was largerthan in the case of hafnium in the two studied solutions. The average drop inthe OCP (ΔEavg) due to fretting was found higher for CP titanium, being morepronounced in 25% serum than in the saline solution. The inital depassiva-tion peak (ΔEmax) was more acute in the case of CP titanium in both solutions.Furthermore, the depassivation peak was more pronounced in 25% serum inhafnium and CP titanium.


The evolution of the current measured in CP titanium and hafnium at a con-stant passive potential (0 V vs Ag/AgCl) is shown in Figure C.4. The test canbe divided in three phases: first from A to B (in Figure C.4), samples were im-mersed 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. Afterthat, from B to C (in Figure C.4), the onset of the fretting led to removal of the


Table C.4: Anodic peak current (Ipeak) and average current (Iavg) measuredduring fretting at a potential of 0 V (vs Ag/AgCl) for Hf and CP Ti.

Ipeak (µA) Iavg (µA)

Hf 0.9% NaCl 6.01 ± 3.24 0.86 ± 0.2525% serum 4.35 ± 2.15 1.03 ± 0.22

CP Ti 0.9% NaCl 14.42 ± 6.16 2.82 ± 0.5525% serum 23.67 ± 6.35 3.57 ± 0.52


passive film, causing a sudden increase in the current due to the exposition ofthe unprotected metal to the media. After an initial peak, the current decreasedtowards lower values due to the equilibrium between mechanical repassivationand electrochemical depassivation. Finally, from C to D (in Figure C.4), thelow 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 TableC.4. The peak currents measured in the case of CP titanium were higher com-pared to hafnium, indicating a more severe initial depassivation. The averagecurrents 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, in-creasing after this initial stage in the case of CP titanium. This could have beencaused by a higher wear and an increased depassivated area.

C.5 Discussion

The behaviour of hafnium and CP titanium in fretting corrosion conditionshas been studied in order to assess the suitability of hafnium for biomedicalapplications.

Previous studies have shown a tendency of hafnium to suffer from localisedcorrosion in simulated body fluids [13]. The effect of fretting on the break-down potential of hafnium was studied by means of cyclic polarisation scans.The breakdown potential observed when hafnium was subjected to fretting wasdrastically 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 corro-sion 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 studiedsolutions (pH 7.4) [14], which would lead to attractive forces towards the an-odically polarised surface. Negatively charged proteins and chlorides, whichare known to be a cause of localised corrosion [15], would compete to ad-sorb on the surface, retarding the formation of pits [16]. The currents reacheddue to the propagation of the pits were reduced when proteins were present inthe media. A similar effect of the organic species has been reported in staticconditions in previous studies [13].

Open circuit potential (OCP) is commonly used as a qualitative indicationof the corrosion regime in which a material resides [17]. Both CP titanium andhafnium were in the passive state when immersed in saline and serum solu-tions. Both materials showed an anodic shift before the onset of the frettingthat suggested thickening of the passive film. When fretting started a suddencathodic shift was observed. The mechanical damage on the surface exposedthe bulk metal to the corrosiveness of the media; therefore the OCP value wasmuch 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. Aninitial peak in the cathodic direction was observed with the onset of the fret-ting. The initial peak was more prominent in the case of CP titanium, whichsuggested a more severe mechanical depassivation when fretting started. Afterthe initial cathodic peak the OCP increased towards more anodic values due tothe balance between the mechanical depassivation due to fretting and the elec-trochemical repassivation or reformation of the oxide layer due to the reactionof the materials with the oxygen in the solution [19]. The difference betweenthe OCP value on the passive state and under fretting was smaller in the caseof hafnium, suggesting less mechanical damage, which could be related to thehigher hardness of the material. A similar trend was observed in both the salinesolution and the serum solution.

The evolution of the anodic current during the tribological tests under pas-sive potential (0 V vs Ag/AgCl) showed higher corrosion currents for CP ti-tanium under the studied conditions. Current increased drastically when themechanical damage started. The initial peak was higher for CP titanium inboth 0.9% NaCl and 25% serum solutions, which could be caused by a moresevere initial depassivation due to the lower hardness of the material. After theinitial peak, lower and more stable current values were reached. The average


currents measured for hafnium were lower than those for CP titanium in thesaline 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 solu-tions, an increase of the current was observed during the test. Proteins couldhave had an effect on the passivation of the materials by restricting the dif-fusion of oxygen to the surface and making repassivation of the metal moredifficult [14]. The negative effect of proteins on the repassivation of the metalwould 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 ofCP titanium, the lower surface hardness might lead to increased mechanicaldamage and therefore more severe mechanical damage.

C.6 Conclusions

Metals used for medical implants can be subjected to fretting corrosion, whichcan compromise the life time of the implant. The fretting corrosion behaviourof hafnium and CP titanium has been investigated. The main conclusions ofthis study are:

• Hafnium tended to suffer from localised corrosion; its resistance wasdrastically decreased when the material was subjected to fretting. How-ever, the organic species contained in the serum solution retarded theformation of pits and limited their propagation.

• Hafnium showed a passive state in both 0.9% NaCl and 25% serum so-lution under static conditions. The passive layer was damaged due tofretting; however, the mechanical depassivation was more severe in thecase of CP titanium in comparison to hafnium. Both hafnium and CPtitanium repassivated in the studied solutions when the mechanical dam-aging stopped.

• The damage produced due to fretting produced increased currents andtherefore higher corrosion of the materials. Hafnium showed lower an-odic 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

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Investigation of Hafnium for Biomedical Applications

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