Corrosion Fatigue of Biomedical Metallic Alloys

Acta Biomaterialia 8 (2012) 937–962

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Acta Biomaterialia

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Review Corrosion fatigue of biomedical metallic alloys: Mechanisms and mitigation ⇑ Renato Altobelli Antunes a, , Mara Cristina Lopes de Oliveira b a Engineering, Modeling and Applied Social Sciences Center (CECS), Federal University of ABC (UFABC), 09210-170 Santo André, SP, Brazil b Electrocell Ind. Com. Equip. Elet. LTDA, Technology, Entrepreneurship and Innovation Center (CIETEC), 05508-000 São Paulo, SP, Brazil article info abstract

Article history: Cyclic stresses are often related to the premature mechanical failure of metallic biomaterials. The com- Received 15 April 2011 plex interaction between fatigue and corrosion in the physiological environment has been subject of Received in revised form 7 August 2011 many investigations. In this context, microstructure, heat treatments, plastic deformation, surface finish- Accepted 9 September 2011 ing and coatings have decisive influence on the mechanisms of fatigue crack nucleation and growth. Fur- Available online 16 September 2011 thermore, wear is frequently present and contributes to the process. However, despite all the effort at elucidating the mechanisms that govern corrosion fatigue of biomedical alloys, failures continue to occur. Keywords: This work reviews the literature on corrosion-fatigue-related phenomena of Ti alloys, surgical stainless Corrosion fatigue steels, Co–Cr–Mo and Mg alloys. The aim was to discuss the correlation between structural and surface Titanium alloy Stainless steel aspects of these materials and the onset of fatigue in the highly saline environment of the human body. Cobalt alloy By understanding such correlation, mitigation of corrosion fatigue failure may be achieved in a reliable Magnesium scientific-based manner. Different mitigation methods are also reviewed and discussed throughout the text. It is intended that the information condensed in this article should be a valuable tool in the development of increasingly successful designs against the corrosion fatigue of metallic implants. Ó 2011 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

1. Introduction size, leading to fracture. This process is known as corrosion fatigue, denoting the failure of a material under the simultaneous action of The selection of a metallic biomaterial to be employed as a load- cyclic loads and chemical attack [9]. The reduction in fatigue life of bearing orthopedic device should be based on a reliable analysis of metallic implants under corrosion fatigue has been well docu- relevant materials properties. Fatigue resistance is perhaps the mented [10–12]. most important issue to be addressed in this case. Several reports Teoh [13] emphasizes the importance of knowing the surface indicate that fatigue-related mechanisms are responsible for the substructure of biomaterials in order to understand the mecha- most part of mechanical failures of implantable medical metallic nisms of fatigue failure. The substructure is composed of three dis- components [1–4]. Chao and López [5] reported that nearly 90% tinct layers – a molecular absorbed layer, the passive oxide film of the surface fracture of cementless hip prosthesis manufactured and the deformed layer – as shown in Fig. 1. Cyclic loadings lead with Ti–6Al–4V alloy was due to fatigue mechanisms. In addition to the generation of wear debris (contact body). The molecular ab- to oscillating mechanical loads, implants are exposed to the phys- sorbed layer consists of growing tissue (cells) in contact with the iological fluid that consists of a saline solution [6] including Na+, physiological environment and the passive layer on the surface of 2+ 2 2 Mg ,Cl ,SO4 and HCO3 . Metallic implants owe their corrosion the metallic implant. The deformed layer arises from the cyclic resistance to the formation of a stable, compact and continuous loadings that cause localized plastic deformation, forming the oxide surface film called passive film that prevents the underlying damage zone in the microstructure of the metallic implant. Many bare metal surface from coming into contact with these aggressive questions arise from this picture: how does the intrinsic micro- ions [7]. However, the passive film may be locally dissolved, espe- structure of the alloy influence the ability of the material to cially by chloride ions, generating pits that rapidly propagate, lead- withstand cyclic loadings without fatigue failure? How does the ing to pitting corrosion. Nucleation of fatigue cracks has been stability of the passive film in contact with the physiological envi- related to the presence of pits on the surface of metallic materials ronment affect corrosion fatigue mechanisms? Does wear partici- [8]. Under fatigue conditions, the aqueous environment can accel- pate in the overall fatigue resistance of the material? How does erate the initiation of a surface flaw and propagate it to a critical one evaluate these phenomena? How does one maximize the performance of the medical device against corrosion fatigue? Answer- ing these questions correctly may be the difference between a ⇑ Corresponding author. Tel./fax: +55 11 4996 3166. successful implant life and a premature catastrophic failure. If E-mail address: [email protected] (R.A. Antunes). one thinks that all these phenomena are self-related and thus have

1742-7061/$ - see front matter Ó 2011 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.actbio.2011.09.012 938 R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962

Fig. 1. Schematic representation of the substructure of a metallic biomaterial (from Teoh [13]; reprinted with permission from Elsevier). a simultaneous action within the human body, the complexity of this task appears to be extreme. A sound knowledge of materials science concepts owing to processing, heat treatment, microstructure of specific metallic alloys, corrosion, wear, fatigue and surface modification methods is needed to achieve a reliable finished product. These issues are explored in this work, with particular emphasis on the more common biomedical metallic materials, i.e. titanium and titanium alloys, austenitic stainless steels, Co–Cr–Mo and biodegradable magnesium alloys.

2. Corrosion fatigue of metallic implants

2.1. Fatigue: basic concepts

The fatigue of metals has been extensively studied [14–17]. The fluctuating stresses typical of fatigue may lead to the initiation and growth of a crack that, upon reaching a critical size, leads to fracture [18]. Fatigue crack initiation is frequently reported to occur at stress concentration sites and manufacturing defects of the metallic component, such as holes, fillets, welds, notches, pits and surface imperfections from machining operations [19–21].If such design or manufacturing factors are absent, fatigue cracks may initiate in microstructural defects, such as grain boundaries and non-metallic inclusions [22–24]. The role of the microstructure in fatigue crack initiation has Fig. 2. TEM image of two slip bands in a failed fatigue specimen of Ni-based been reviewed recently by Chan [25]. The term ‘‘crack initiation’’ superalloy (Udimet 720). The inset is a low-magnification view of persistent slip may be defined from a scientific or from an engineering point of bands (from Sangid et al. [31]; reprinted with permission from Elsevier). view. The scientific approach states that initiation is the number of cycles required to generate, nucleate or form the smallest A transmission electron microscopy (TEM) image of two slip detectable crack. Chan adapted this definition to mean the process bands in a fatigue-failed specimen of an Ni-based superalloy (Udi- of forming a fatigue crack of a length that is of the same order of a met 720) is shown in Fig. 2. The inset shows a low-magnification grain size or less. The engineering definition is based on the small- view of the PSBs in this region. Similar dislocation arrangements est crack that can be determined by reliable non-destructive eval- of PSBs also form in polycrystalline ductile metallic alloys. How- uation techniques, such as X-ray and ultrasonic methods [26,27]. ever, in this case, the interaction between PSBs and grain bound- Independently of the definition, it is widely accepted that fatigue aries or twin boundaries make the study of fatigue crack crack initiation in ductile metals is the result of localized plastic initiation mechanisms more complex than in single crystals deformation during cyclic slip processes [28–30]. Plastic strain [32,33]. Furthermore, precipitates, non-metallic inclusions and can be highly concentrated in some regions while the surroundings dependence on grain crystallographic orientation (texture) also areas are still in the domain of elasticity. The plastically deformed affect the PSB morphology [34]. regions are referred to as persistent slip bands (PSBs). This process Extrusions and intrusions are formed within PSBs. These regions is typical of single crystals of face-centered cubic (fcc) metals. A are preferential sites for fatigue crack initiation. Extrusions and PSB consists of a large number of slip planes, forming a flat lamellar intrusions within a PSB are schematically shown in Fig. 3a. A scan- structure divided into channels by a periodic array of dislocations. ning electron microscopy (SEM) image of an extrusion formed A complete description of PSBs structure and formation is given by within a PSB in high-purity polycrystalline nickel is shown in Suresh [9]. Fig. 3b. R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962 939

Fig. 3. (a) Schematic representation of extrusions and intrusions formed within a PSB as a result of cyclic strain during fatigue of a metallic material (from Schijve [35]; reprinted with permission from Elsevier); (b) SEM image of extrusions within a PSB of high-purity polycrystalline nickel (From Weidner et al. [29]; reprinted with permission from Elsevier).

Fatigue cracks initiated by the intrusion–extrusion mechanism In this virtual infinite life design, fatigue data are presented in the from PSBs are called Stage I cracks. As these cracks propagate form of Wöhler’s or S–N curves, which represent the plot of stress, and grow they become Stage II fatigue cracks. In stage I fatigue S, vs. number of cycles to failure, N [39]. These curves are deter- crack growth the crack and the zone of plastic deformation sur- mined for high cycle fatigue conditions, i.e. for N >105 cycles. The rounding its tip are restricted to only a few grain diameters. This stress level is relatively low and the overall strains are mainly elas- is called the short crack propagation stage [36]. As the stresses in- tic, even though the material undergoes localized plastic deforma- crease, the plastic zone at the crack tip involves a large number of tion. This is a common approach to the evaluation of fatigue grains. This stage leads to the formation of the so-called fatigue resistance of metallic implants [40–42]. striations that appear as ripples on the fracture surface [9,36].An A schematic representation of an S–N curve is shown in Fig. 5.In example of fatigue striations on the fracture surface of a surgical this figure, the cyclic stress range (Dr = rmax rmin) is plotted 316L stainless steel plate is shown in Fig. 4. At this point, it is inter- against the number of cycles to failure, N, in log–log scales. The esting to define the fatigue crack growth mechanism in a more endurance, or fatigue limit (rft), is defined as the stress range be- mechanistic basis. The reliable prediction of crack growth under low which there is no crack growth (the material would be sub- cyclic loadings is a valuable and challenging topic for actual engi- jected to an infinite number of stress cycles without fatigue neering applications of metallic materials. Successful design fracture). The number of cycles corresponding to the endurance against fatigue is often determined by the understanding of this limit is often considered to be 107; this represents an infinite fati- topic. gue life. However, some materials do not present an endurance In this context, it is important to describe how the fatigue life of limit. For these materials, the stress continuously decreases with a metallic component can be evaluated. The prediction of fatigue increasing number of cycles, and the ‘‘fatigue limit’’ is arbitrarily life may follow two different approaches: cumulative fatigue defined as the fatigue strength at a specific number of cycles, typ- damage (CFD) and fatigue crack propagation (FCP) [38]. CFD is re- ically 108 or 5 108 cycles [39]. A distinct fatigue limit has been lated to total fatigue life, safe life or damage-intolerant life design. reported for titanium alloys, Co–Cr–Mo and austenitic stainless

Fig. 4. Fatigue striations on the fracture surface of a surgical 316L plate (from Fig. 5. Schematic representation of an S–N curve (from Plekhov et al. [43]; Kanchanomai et al. [37]; reprinted with permission from Elsevier). reprinted with permission from Elsevier). 940 R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962 steel implants [44–46], while biodegradable magnesium alloys do not exhibit an endurance limit [47]. For low numbers of cycles (low cycle fatigue; N <105 cycles), cyclic stresses are higher than for high cycle fatigue. The resulting plastic deformation makes cyclic strain more relevant than cyclic stress. Hence, fatigue tests are controlled by cyclic strain instead of cyclic stress in the low cycle fatigue regime, plotting plastic strain range (Dep) vs. the number of cycles to failure. Despite the higher number of investigations on the high cycle fatigue approach of biomaterials, low cycle fatigue data have been reported by some authors [48,49]. This approach is important in applications such as pacemakers or clasps for removable partial dentures [50]. The cumulative fatigue design approach has well-established limitations. As pointed out by Zhang et al. [51], the accuracy of life prediction is strongly influenced by the geometry of the component under testing (stress concentration factors), process and load, resulting in a wide scattering of fatigue life data. Furthermore, it ignores the presence of flaws in the material [39]. The fatigue crack propagation (FCP) approach overcomes these drawbacks. This method derives from the linear elastic fracture mechanics (LEFM) Fig. 6. Schematic representation of a typical da/dN vs. DK plot (Paris’s law curve) concepts developed in the early 1960s, and is also known as dam- (from Jones et al. [57]; reprinted with permission from Elsevier). age-tolerant or finite life design. In their pioneering work, Paris et al. [52] proposed that the stress intensity factor range

(DK = Kmax Kmin) could be used to characterize the rate of crack this curve. Region I corresponds to the regime in which the average advance per cycle (da/dN) during fatigue. The plot of da/dN vs. crack growth increment per cycle is smaller than a lattice spacing DK allows the interpretation of experimental fatigue data in terms [9]. A threshold value of the stress intensity factor range (DKth)is of the well-known power-law relationship called Paris’s law: marked on the plot, and determines the end of region I and the start of region II. According to Carpinteri and Paggi [38], this da ¼ C:DKm ð1Þ parameter is closely related to the concept of fatigue limit. The dN conventional definition of DKth is taken as the value of DK below In Eq. (1), C and m are the Paris coefficient and exponent, which the crack growth rate is 109 mm per cycle or less. Paris’s respectively, which depend on the material, environment, law is valid in region II, which is related to the propagation of long frequency and stress ratio (R = rmin/rmax). DK is expressed as: cracks (stage II cracks). Region III denotes fast crack propagation, ffiffiffiffiffiffiffiffi p leading to final failure at very high values of DK, approaching KIC. DK ¼ Y:Dr: p:a ð2Þ KIC is a particular property of the material, called the fracture In Eq. (2), Y is a dimensionless parameter that depends on the toughness, and is related to its resistance to brittle fracture. geometry and size of an existing crack of length a in the material. FCP philosophy is used by several authors to investigate the At this point, it is important to emphasize that the LEFM approach crack growth rate of metallic implant materials [58–60]. However, only provides reliable data for the fatigue crack propagation rate of despite the widespread use of the FCP approach, several deviations metallic materials when the initial size of a fatigue crack is longer from the behavior described by Paris’s law have led to the develop- than 1 or 2 mm and the small-scale yielding conditions apply [9]. ment of other fatigue crack growth criteria. It is not within the These fatigue flaws are classified as long cracks. Conversely, the scope of this article to give a deep description of these multiple cri- growth rates of small cracks can be much faster at the same DK. teria. For a more complete insight, the reader should consult the Suresh and Ritchie [53] gave a detailed definition of small work by Pugno et al. [61] and references therein. Despite the lim- cracks: (i) microstructurally small cracks if the crack size is compa- itations of both the CFD and FCP approaches, the general bases of rable to the scale of the characteristic microstructural dimension, these design philosophies are perfectly sufficient for the discussion such as the grain size for monolithic materials; (ii) mechanically of the literature on corrosion fatigue phenomena of metallic im- small cracks if the plastic zone around a crack tip is comparable plants that is presented in the following sections. to the crack size; (iii) physically small cracks if the crack is larger than the characteristic microstructural dimension and the scale 2.2. Corrosion-assisted fatigue of metals: general mechanisms of the plastic zone and is smaller than 1 or 2 mm; and (iv) chemi- cally small cracks when environmental stress corrosion mecha- The classification of corrosion-assisted fatigue of metals was nisms become more active below a determined crack size. For primarily given by McEvily and Wei [62]. The environment affects these cracks, the use of Eq. (1) is no longer applicable. fatigue crack growth depending on the Kmax value applied to the The behavior of small fatigue cracks is of prime importance, KISCC threshold for a sustained load. Kmax is the stress intensity fac- since a major part of the fatigue life of a metallic component is tor and KISCC is the stress corrosion cracking (SCC) threshold. The spent in the propagation of this type of crack. Since the LEFM ap- classification scheme is separated into three different possibilities proach is not valid in this case, the propagation behavior of small according to da/dN vs. Kmax plots, as shown in Fig. 7. Type A behav- fatigue cracks has been described by other criteria. The is invited ior is related to the situations where Kmax,th is reduced by the envi- to consult the reports by Pearson [54], Endo and McEvily [55] ronment. For high values of da/dN, the curves of aggressive and and Shyam et al. [56] provide deeper insights into this theme. In inert environments are superimposed, that is, the contribution of spite of the limitations of the LEFM approach to long fatigue cracks, the environment to fatigue crack growth is reduced with increas- the fatigue propagation rate behavior of biomedical metallic alloys ing fatigue crack growth rates. It is deduced, then, that corrosion has been widely characterized through da/dN vs. DK plots. fatigue is a time-dependent process and that such dependency be- A typical da/dN vs. DK plot has a sigmoidal shape, as schemat- comes less pronounced with increasing frequency of cyclic stress. ically shown in Fig. 6. Three distinct regions may be defined from This is considered to be a true corrosion fatigue behavior as the R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962 941

Fig. 7. Classiﬁcation scheme of corrosion-assisted fatigue crack growth proposed by McEvily and Wei: solid line refers to aggressive environment; dotted line refers to inert environment (from Vasudevan and Sadananda [63]; reprinted with permission from Elsevier).

environment affects fatigue crack growth even when Kmax < KISCC. The influence of the environment is less marked for type B behavior than for type A at low crack growth rates (Kmax < KISCC). When Kmax exceeds KISCC there is a superposition of the stress corrosion process on fatigue. A plateau is then formed, which is typical of the stress corrosion cracking growth mechanism. Crack growth is governed mainly by the applied stress. Type C behavior is a combination of types A and B. Crack growth is dependent on both time (frequency of the cyclic stress) and the applied stress. Recently, Vasudevan and Sadananda [63] extended this early classification scheme of corrosion-assisted fatigue growth to five types of behavior. The scheme is based on the parameters DK and Kmax that are recognized to provide two crack tip driving forces simultaneously required for the propagation of fatigue cracks. In this way, the corresponding threshold values of DK and Kmax must be surpassed to allow the growth of a fatigue crack. These thresh- ⁄ olds are represented as DK and Kmax. Crack growth is controlled by ⁄ cyclic amplitude in the Paris regime. In this case DK = Kmax and the ⁄ plot of DK vs. Kmax for different crack growth rates (da/dN) will be a straight line. This is characteristic of a pure fatigue crack growth or ideal fatigue behavior, and may be observed for alloys tested in a very high vacuum. A representation of such a mechanism is shown Fig. 8. Representation of an ideal fatigue behavior in trajectory map (plot of DK⁄ vs. K ). In this curve da/dN 2>da/dN 1 (from Sadananda et al. [64]; reprinted with in Fig. 8. The curve in Fig. 8 may be considered as a trajectory of max permission from Elsevier). crack growth mechanism and is known as a trajectory map [64]. Alterations of the mechanism of fatigue crack growth give rise to deviations from this ideal behavior. It has been reported that all behavior typical of materials that are insensitive to the environ- deviations from the ideal behavior occur when Kmax exceeds DK⁄, ment. Type II is the true corrosion fatigue behavior as observed ⁄ i.e. non-ideal behaviors fall below the line DK = Kmax, approaching in Fig. 7a. Type III behavior accounts for the case where the envi- the x-axis [64]. One main reason for such a deviation is the onset of ronmental fatigue crack growth curve in a da/dN vs. DK plot runs a corrosion process. According to Vasudevan and Sadananda [63],if parallel to the pure fatigue behavior. The contribution of the envi- a metallic material is in a corrosive environment, Kmax controls ronment is nearly constant throughout the range of fatigue crack crack propagation by breaking the crack tip bonds. One possibility growth. Type IV is typical of a superimposed stress corrosion is that the environment reduces the Kmax required to break the behavior on fatigue analogous to the type B behavior proposed bonds. The trajectory map is influenced by the aggressiveness of by McEvily and Wei [62] (Fig. 7). Type V behavior is an extreme ⁄ the environment. As the environmental effect increases relative case of Type IV. The slope of DK vs. Kmax tends to zero as the tra- to fatigue, the crack growth trajectory can become almost parallel jectory crack growth is parallel to the Kmax axis. The trajectory to the x-axis. The trajectory map may thus be used to study the maps for these five types of corrosion-assisted fatigue crack crack growth mechanisms for a given material–environment pair. growth are shown in Fig. 9. This approach provides an effective The classification scheme proposed by Vasudevan and Sadananda method of separating the relative contributions of fatigue and [63] can be described from these definitions. environment to the crack growth of specific materials. Despite Type I is related to the case where fatigue crack growth is con- the valuable information of trajectory maps to the investigation ⁄ trolled by cyclic amplitude and DK = Kmax, i.e. it is the ideal fatigue of corrosion fatigue of metallic materials, we have not encountered 942 R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962

Fig. 9. Trajectory maps typical of the five types of fatigue crack growth behavior according to the classification scheme of Vasudevan and Sadananda [63]. Cleavage and ⁄ microvoid coalescence (monotonic deformation modes) can produce a sharp decline in the DK vs. Kmax curve at high values of Kmax (from Vasudevan and Sadananda [63]; reprinted with permission from Elsevier). any reference in the literature that refers to such an approach con- harmful effects that it exerts on the mechanical stability of the pas- cerning the fatigue behavior of biomedical metallic alloys. Hence, sive films. Fretting corrosion is a crucial mechanism of in vivo im- this is an open field of research in the biomedical area. As depicted plant corrosion. Crevice corrosion is frequently associated with this by Vasudevan and Sadananda [63], the understanding of the gov- phenomenon [82,83]. Rabbe et al. [84] confirmed the validity of erning mechanisms of fatigue crack growth through the applica- this model for Ti6Al4V and 316L biomedical alloys. Chandraseka- tion of trajectory maps can help develop more reliable fatigue ran et al. [85] studied the fretting corrosion behavior of Ti–6Al– life prediction methods. 4V modular taper specimens and discussed the clinical implica- In the next sections, we give, wherever possible, a classification tions of this phenomenon. The potential adverse effects of fretting of the type of corrosion fatigue behavior of different biomedical damage are the loosening of modular connections and the forma- metallic materials according to the published data of fatigue crack tion of debris that can lead to osteolysis or accelerate articular growth (da/dN vs. DK curves) in different environments. wear if entrapped between the contacting surfaces. It is noticeable that fretting motion between modular components in orthopedic 2.3. The role of wear implants is typically enhanced by manufacturing-induced discontinuities, such as scratches introduced during machining opera- Tribocorrosion is defined as an irreversible transformation of a tions. Hence, the careful control of the surface integrity of the material from concomitant physicochemical and mechanical sur- components is highly recommended in order to diminish the fret- face interactions occurring at tribological contacts [65]. The impor- ting corrosion of biomedical alloys. tance of this phenomenon for metallic implants has been widely Barril et al. [86] identified a clear correlation between mechan- recognized [66–68]. This interest arises from the widespread use ical and electrochemical parameters by evaluating the fretting cor- of total hip arthroplasty as a common surgical intervention. A rosion behavior of a Ti–6Al–4V plate against an alumina ball. They sliding contact is formed between the femoral head and the ace- showed that the electrochemical activity depends on the third- tabular cup during the articulating movements of the hip joints body behavior in the contact. Current peaks were associated with in the human body [69]. The sliding surfaces are then subjected third-body accumulation in the contact, suggesting that the to tribocorrosion. The resulting particulate debris has been found third-body particle formation occurred due to local depassivation to promote adverse biological effects [70,71]. Furthermore, the of the metal. The depassivated surface then dissolves and repassi- small relative displacements between the prosthesis and the sur- vates. The freshly formed metallic wear particles may also dissolve rounding bone tissues lead to deterioration of the biomedical al- and passivate, contributing to the observed current peaks. loys by fretting corrosion [72]. Documented cases of fretting Diomidis et al. [87] also observed the repassivation of biomed- corrosion have been reported in bone plates and screws at the ical titanium-based alloys during fretting corrosion experiments. bone–stem interface and stem–cement interfaces of modular hip Monticelli et al. [88] reported that the corrosion processes of Ti– implants [73]. Titanium alloys are susceptible to fretting damage 6Al–4V were stimulated markedly when the material was sub- [74,75]. In the same way, fretting corrosion has been a concern jected to combined wear and corrosion conditions, while under for austenitic stainless steels and cobalt-based alloys [76–78]. pure corrosion the corrosion current density was two orders of Two- or three-body contacts are frequently associated with magnitude lower. tribocorrosion [79]. Entrapped wear debris acts as an abrasive Recently, Souza et al. [89] brought new knowledge to the field and is defined as the third body, while the sliding surfaces are by investigating the influence of oral biofilms on the biotribocorro- the first and second bodies [80]. The main concerns while one sion behavior of commercially pure titanium (grade 2). The tests investigates the simultaneous action of corrosion and wear in were conducted in vitro by exposing the samples to an artificial biomedical systems are the ability of the passive layer to withstand saliva solution. The authors reported that the titanium surfaces the mechanical stresses arising from wear, the ability of the metal can be easily colonized by oral biofilms. The presence of biofilms surface to repassivate when the passive film is removed and the decreased the corrosion potential of the titanium implants proba- resistance of the new repassivated surface to both wear and corro- bly due to the release of acidic substances that diminished the pH sion [81]. In this regard, fretting has a big influence on the corro- of the test solution in comparison with a condition without the sion behavior of orthopedic devices due to the potentially presence of a biofilm on the surface of the samples. During the R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962 943 tribocorrosion tests in the absence of a biofilm, the potential carbides is suppressed and wear becomes less intense for a high- dropped due to the removal of the TiO2 passive film under an ap- carbon CoCrMo. However, this effect was not observed in the plied normal load of 200 mN load. Upon unloading, the potential PBS–albumin solution. The authors recognize the complexity of increased due to the repassivation of the titanium surface. The the interaction between the electrolyte solution and the tribocor- potential remained almost unchanged after unloading for the sam- rosion behavior of CoCrMo alloys, and the necessity for further ples tested at a normal load of 100 mN, suggesting that repassiva- investigations to clarify the mechanisms involved in this process. tion could occur during sliding. The presence of biofilms altered In addition to the tribological and chemical aspects depicted this behavior. The potential drop after starting the sliding tests at above, the concomitant action of cyclic stresses should not be dis- 200 mN was also verified, but at longer times than when no bio- regarded. Fretting fatigue develops under the combination of wear films were present. Hence, it is suggested that the biofilm is first and fatigue processes when two contacting surfaces experience removed and then the TiO2 passive film. The authors noticed that small oscillatory relative motions while enduring a normal load the corresponding coefficient of friction was lower during the time and at least one of them carries bulk stress [99,100]. In this regard, the biofilm remained on the surface of the samples. When the slid- Hoeppner and Chandrasekaran [101] reviewed the fretting-related ing tests were conducted at 100 mN, the potential remained un- damage of orthopedic implants. Fretting causes a local increase in changed, as did the coefficient of friction, which remained low strain on the surface of the metallic material that is a preferable throughout the whole test. The wear scars produced on the surface site for the nucleation and early growth of fatigue cracks, leading of the samples after the tribocorrosion tests were observed by SEM. to premature fatigue failure. Fretting fatigue life is typically di- It was found that the wear scars were larger in the absence of bio- vided in two parts, the first of which is referred to as stage I prop- films, but were also present when a biofilm was present. This result agation. In this stage, the crack initiates along a slip direction in a points to a positive effect of the oral biofilm on the wear resistance grain. The second part is stage II crack propagation, where the fa- of the titanium surface, although the passive film was still removed tigue crack growth can be predicted by the Paris law. Stage II from the surface of the samples covered with a biofilm, especially evolves into fast propagation when the fatigue crack is longer than at high loads. Electrochemical impedance spectroscopy measure- a critical size [100], which was experimentally shown to be 1.5–2.0 ments conducted in the absence of wear showed that biofilms de- times the grain size of the metallic alloy [102]. Waterhouse and creased the corrosion resistance of the material. It is concluded Dutta [103] recognized a long time ago that the corrosive nature that a wear corrosion mechanism is developed during sliding of of the body fluids affects the fretting fatigue behavior of biomedical the titanium samples and this process may lead to the failure of titanium alloys. The importance of considering the interaction of dental-implant-supported systems. biomedical alloys with living tissues for a complete understanding Igual-Muñoz and Julián [90] studied the tribocorrosion behavior of corrosion fatigue phenomena has also long been recognized by of a high-carbon CoCrMo biomedical alloy. The current density was Morita et al. [104]. They observed a remarkable reduction in the fa- found to increase in the passive domain of the alloy. tigue strength of austenitic stainless steel biomedical alloys when According to Mischler [91], the corrosion rate of passive metals the fatigue tests were conducted in vivo. Although the authors did can increase by several orders of magnitude under tribocorrosion not evaluate the influence of fretting on the in vivo corrosion fati- conditions in comparison with pure corrosion. There is a constant gue of the biomedical alloys, this phenomenon is likely to be pres- depassivation followed by repassivation during the combined wear ent. Chandra et al. [100] recently studied the fretting corrosion and corrosion processes. Consequently, a sharp increase in the cur- behavior of Ti–6Al–4V. They observed that the wear volume de- rent follows in tribocorrosion tests at the onset of rubbing, thus pends on the stress state of the alloy and on the pH of the electro- accelerating the wear corrosion process of the alloy, leading to pro- lyte in contact with the surface of the metallic material. They also gressive consumption of the material. evaluated the fatigue crack growth under fretting conditions, and Sinnett-Jones et al. [92] showed the influence of the integrity of proposed a model to predict the critical crack length and the life the passive film and of the repassivation kinetics on the tribocorro- expectancy of the component under fretting fatigue. According to sion rate of a CoCrMo alloy. Chandra et al., fretting motion in the transverse direction of the Yan et al. [93] showed that under tribological contact the corro- throat stem in a modular hip joint decreases the crack length that sion rate of high-carbon CoCrMo, low-carbon CoCrMo and 316L is critical for fast crack propagation. Moreover, both the critical stainless steel alloys increased 20–60 times in comparison with crack length and life expectancy diminish with the progression of static corrosion conditions. The constant depassivation of the al- wear. Furthermore, the amount of wear depends on the pH of loys by the mechanical forces generated in the tribocorrosion test the electrolyte. Hence, the corrosive conditions of the fluid in the accelerated the corrosion processes. Similar findings have been re- joint of a modular orthopedic device are strictly related to the crit- ported by the same group in other publications [94,95]. ical crack size under fretting fatigue conditions. Igual-Muñoz and Mischler [96] studied the influence of differ- The tribocorrosion and fretting corrosion mechanisms dis- ent simulated body fluids on the tribocorrosion behavior of cussed to this point are mainly referred to in vitro laboratory inves- CoCrMo biomedical alloys. NaCl and phosphate-buffered saline tigations. However, the in vivo behavior of metallic implants under (PBS) solutions with and without the addition of albumin were combined wear corrosion or fretting corrosion has been hardly used in the tests. The wear ranking of the alloy depended signifi- studied. Lomholt et al. [105] gave a recent contribution towards cantly on the test electrolyte. The composition of the solution the understanding of the in vivo behavior of Ti–6Al–4V biomedical influences the surface chemistry of the metallic material. Conse- alloy under tribocorrosion conditions. They compared results from quently, mechanical wear is also affected. The nature of the passive laboratory tribocorrosion tests with those from retrieved acetabu- film depends on the electrolyte, and can alter the mechanical lar cups with appurtenant polymer liners from patients that suf- response of metals in sliding tribocorrosion systems [97]. The fered from pain. The implants were part of artificial hip joints. adsorption of ions or molecules from the electrolyte may affect SEM images of the retrieved implants and the laboratory-tested the repulsion/attraction forces between wear particles and solid specimens evidenced similar features. A pattern of light and dark surfaces, modifying the response of the material to wear. The contrast was found on the failure surfaces of the retrieved implants presence of proteins in the electrolyte solution during tribocorro- and the specimens evaluated in vitro. Higher microhardness values sion test has been reported to decrease wear-induced corrosion were found in the dark regions that also presented cracks for both currents of CoCrMo alloys [98]. Indeed, Igual-Muñoz and Mischler conditions (in vivo and in vitro). Metallic particles were found to be [96] suggested that in the NaCl–albumin solution the pulling out of adhered to the polymer liners of the retrieved implants and on the 944 R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962 electrolyte solution used in the tribocorrosion tests. A mechanism behavior of biomedical titanium alloys based on either the CFD for the failure of the retrieved implants has been proposed based or FCP approach. on the experimental observations. The movement of the femoral head within the acetabular cup after the polymer liner has been 2.4.2. The CFD approach worn through generates wear debris from the acetabular cup. Akahori et al. [118] investigated the fatigue and corrosion fati- The accumulation of the wear debris from the Ti–6Al–4V acetabu- gue behavior of Ti–29Nb–13Ta–4.6Zr (TNTZ) alloy in air at 22 °C lar cup forms large flakes. These flakes are then pressed into the and in Ringer’s solution at 37 °C. The material was tested under dif- surface during the sliding between the femoral head and the ace- ferent conditions: (a) solutionized at 1063 K for 3600 s in argon tabular cup. The dark contrast on the failure surface of the acetab- atmosphere; (b) the same treatment described in (a) followed by ular cup would arise from this mechanism. As these dark regions cold rolling up to a reduction ratio of 87.5% at room temperature; are brittle, cracks are formed and yield fine particulate debris upon (c) same treatment described in (a) followed by water quenching. further sliding in the hip joint. The transport of the debris through The specimens were then aged at different temperatures, followed the body can then cause inflammation of the peri-prosthetic tissue by water quenching. The main goal of this procedure was to obtain and loosening of the implant. Moreover, the debris also acts as distinct microstructures and to evaluate their effect on the fatigue three-body wear particles, increasing the deterioration of the slid- and corrosion behavior of the alloy. The specimens were tested un- ing surfaces. Although this mechanism makes a valuable contribu- der plain and fretting fatigue conditions. The authors used only the tion to the understanding of the failure of titanium-based CFD approach, showing their results as S–N curves for each set of acetabular cups under in vivo tribocorrosion conditions, it does test conditions. They found that the fatigue strength of both the not encompasses the influence of fatigue on the failure process. solutionized and aged specimens was equal in air and in Ringer’s Indeed, the synergism between wear, fatigue and corrosion of solution. Hence, the contact with the physiological solution did metallic implants is as yet unexplored in the literature. However, not degrade the fatigue behavior of the TNTZ alloy from the stand- this gap in our knowledge has started to be filled by the investiga- point of the cumulative fatigue design approach. However, the fa- tions performed by von der Ohe et al. [106,107]. This group pro- tigue crack growth behavior was not assessed. It is possible that posed a multi-degradation mechanism of passive metals that the immersion in Ringer’s solution can lead to an increase in the combines the simultaneous influence of wear, corrosion and cyclic crack growth rates (da/dN) in comparison with the rates obtained loading. The mechanism is proposed as a model for the behavior of in air. This effect cannot be ignored if the corrosion fatigue behav- metallic structural materials in seawater; however, as it is related ior of the metallic biomaterial is to be fully characterized. to the degradation of passive metals under the combined action of The fatigue behavior of another b alloy, Ti–13Zr–13Nb, was wear, corrosion and fatigue, it could be extended to other environ- characterized by Baptista et al. [119]. They used the CFD approach, ments, such as body fluid. This approach could hence prompt fur- obtaining S–N curves for the alloy in air and in 0.9 wt.% NaCl solu- ther contributions to the understanding of the influence of wear on tion. The alloy was found to be insensitive to the environment. The the corrosion fatigue of metallic implants. curves obtained in air and in the saline solution were equivalent. Similar findings were reported by Akahori et al. [117] for TNTZ b 2.4. Titanium alloys alloy in air and in physiological solution, as depicted above. Boehl- ert et al. [120] have also demonstrated such insensitivity for Ti– 2.4.1. General aspects 15Al–33Nb and Ti–21Al–29Nb alloys, the microstructures of which Pure titanium has two different allotropic phases: below 883 °C consisted mainly of the body-centered cubic b phase. Based on its crystalline structure is hexagonal close-packed (a phase), while these results, Majumdar et al. [121] disregarded the evaluation of above this temperature it is body-centered cubic (b phase). the fatigue behavior of Ti–13Nb–13Zr and Ti–13Nb–13Zr–0.5B in Depending on the alloying elements and the heat treatment used, simulated body fluids, performing the fatigue tests only in air. the final microstructure can be conveniently tailored, according to Azevedo [10] investigated the failure of a commercially pure the desired mechanical products. Titanium alloys are classified as titanium (grade I) reconstruction plate for osteosynthesis. The a, near-a, a + b and b alloys regarding their chemical composition, microstructure of the material consisted of equiaxed a-grains the content and nature of alloying elements and the resultant and intergranular b-platelets. Failure occurred as a consequence microstructures [108]. Aluminum is an a-stabilizer element. Vana- of intergranular corrosion attack on the b-platelets due to the con- dium, niobium, molybdenum and tantalum are b-stabilizers. Zirco- tact with body fluids and fatigue caused by high stress concentra- nium is considered a neutral element [109]. Alloys that consist of a tions at lateral notches in the titanium plate. Azevedo emphasized large fraction of the less ductile a phase, such as the widely used that the loosening of the plate-screws in the bone due to a techni- biomedical alloy Ti–6Al–4V, have a modulus of elasticity (E) cal error of the surgeon or bone resorption also contributed to the around 110 GPa. This value is significantly higher than that of a failure. In another report, Azevedo and dos Santos [122] showed typical human bone (30 GPa) [110]. Such a stiffness mismatch is that the number of cycles to failure of a commercially pure tita- associated with an insufficient load transfer from the metallic im- nium plate (grade I) was reduced during laboratory fatigue testing plant to the bone tissue that may lead to the loosening of the pros- when the tests were conducted under immersion in serum at 37 °C. thesis. This phenomenon is referred to as the stress shielding effect Papakyriacou et al. [123] used the CFD approach to study the fati- [111]. In order to avoid such complications, the development of gue behavior of Ti–6Al–7Nb alloy and commercially pure titanium new biomedical titanium alloys was concentrated on b alloys that in air and in 0.9 wt.% NaCl solution. The number of cycles to failure are typically less stiff than a and a + b alloys (modulus of elasticity was found to decrease in the saline solution for both materials. as low as 50 GPa [112]). Zavanelli et al. [124] found a similar trend for Ti–6Al–4V and com- The fatigue behavior of biomedical titanium alloys has been mercially pure titanium dental implants when exposed to syn- found to be dependent on microstructure [113–115]. A wide vari- thetic saliva and fluoride synthetic saliva. ety of microstructures can be produced depending on the combination of mechanical processing and heat treatment [116,117]. The 2.4.3. FCP approach fatigue and corrosion fatigue behaviors of titanium alloys vary Bache and Evans [125] studied the fatigue crack growth behav- accordingly. Representative data taken from selected references ior of Ti–6Al–4V in air and in 3.5 wt.% NaCl solution at room tem- are referred to in this section. The aim is to provide a clear corre- perature. This alloy is typically used as an implant material, and its lation between microstructural features and corrosion fatigue microstructure is classified as a + b. The authors evaluated three R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962 945 distinct microstructures obtained from different heat treatments: increase in crack growth rates when the material is in contact with mill annealed, bimodal a + b and transformed lamellar. These are an aggressive environment. An interesting aspect described in the shown in Fig. 10. The transformed lamellar microstructure is text is the influence of crystal orientation (texture) on the fatigue formed by colonies of aligned, elongated grains separated by a behavior of titanium alloys. When the basal planes of the hexago- continuous retained b phase. It is interesting to note that the mill nal a phase are subjected to a tensile load, corrosion fatigue crack annealed and bimodal alloys were nearly insensitive to the envi- growth depends on the frequency of the cyclic stress. The crack ronment, with only a slight increase in crack growth rates (da/ growth rates increase when the frequency is decreased. If no basal dN) upon immersion in the saline solution in comparison to the planes are subjected to a tensile load, then corrosion fatigue crack rates obtained in air. Conversely, transformed lamellar alloys pre- growth is independent of the frequency of the oscillating stress. sented a marked increase in crack growth rates under immersion Conversely, if the basal planes are subjected to a perpendicular in the saline solution. This result points to a clear and decisive tensile stress, the environmental effects on fatigue crack growth influence of microstructure on the corrosion fatigue behavior of are marked. This result was also reported by Hall [127] for a Ti– Ti–6Al–4V. The fatigue crack propagation approach successfully 6Al–4V alloy. The control of crystallographic texture of plates indicated this influence. According to the classification scheme of and bars can be done through thermo-mechanical processing such corrosion-assisted crack growth proposed by Vasudevan and as rolling and extrusion. Furthermore, the microstructure of the al- Sadananda [63], Ti–6Al–4V with mill annealed and bimodal micro- loy is also important as the crystallographic effect is only related to structures are type I materials, that is, they present ideal fatigue the hexagonal a phase, not the cubic b phase. In this respect, cor- behavior. Transformed lamellar microstructure is a type II material rosion fatigue crack growth of b titanium alloys is insensitive to at low stress ratios (R) and a type IV material for increasing values variations in the frequency of the cyclic stress. For these alloys, fa- of R. The increment of crack growth rates with the increase in the tigue data obtained in air could be used to predict fatigue life. stress ratio is typical of environmentally assisted crack growth. Un- Moreover, low- to medium-strength a and a + b alloys (yield der high R the crack does not ‘‘close’’ during the unloading cycle. As strength less than 700 MPa) behave in the same way as b alloys. a consequence, a constant supply of the corrosive species reaches The corrosion fatigue crack growth behavior of laser melted (LS) the crack tip and the reaction products are transported away. Such and ingot metallurgy (IM) Ti–6Al–4V under a vacuum, in air and in knowledge is essential to the design of biomedical implants against a 3.5 wt.% NaCl solution was studied by Lee et al. [128]. They found fatigue. The most favorable microstructure should be preferred in that crack growth rates are much higher in air and saline solution order to avoid premature failures. than in a vacuum for both LS and IM alloys. It is noteworthy that Gregory [126] reviewed the corrosion fatigue crack growth the behavior in air and in the saline solution is equivalent when behavior of titanium alloys and commented on the expected R = 0.1. Nevertheless, when R = 0.9 the crack growth rate in

Fig. 10. Microstructures of Ti–6Al–4V alloy: (a) mill annealed; (b) bimodal; (c) transformed lamellar (from Bache and Evans [125]; reprinted with permission from Elsevier). 946 R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962

Fig. 11. Microstructures of: (a) LS and (b) IM Ti–6Al–4V alloy (from Lee et al. [128]; reprinted with permission from Elsevier).

3.5 wt.% NaCl increases more rapidly than in air. The authors state that a pure corrosion fatigue behavior operates at R = 0.1. The alloys may then be classified as type II according to the classification scheme of Vasudevan and Sadananda [63].AtR = 0.9, a stress corrosion cracking process would be present in addition to corrosion fatigue, leading to substantial differences between the da/dN curves of the materials exposed to air and in the saline solution. In this case, the corrosion-assisted fatigue crack growth behavior of both materials would be classified as type IV. In addition to the effect of the environment on the fatigue crack growth behavior of the alloys, Lee et al. [128] also assessed the influence of microstructure. The LS alloy microstructure consisted of large b grains (diameter of 600 lm) containing a platelets, while the IM alloy had much smaller, equiaxed a grains (diameter of 15 lm) in a matrix of transformed b, containing a platelets (Fig. 11). The IM alloy was found to present higher fatigue crack growth rates. This behavior was attributed to its large b grains. The authors also expressed their data as trajectory maps (DK⁄ vs. Kmax plots), as shown in Fig. 12 for R = 0.9. It can be seen that the trajectories of both alloys in air and in 3.5 wt.% NaCl solution start ⁄ at lower values of DK and Kmax than in a vacuum. There is a clear divergence from the ideal fatigue behavior even in a vacuum. The environment provides an additional driving force Fig. 12. Trajectory maps of LS and IM Ti–6Al–4 V alloys in different environments from Lee et al. [128]; reprinted with permission from Elsevier). of chemical nature to fatigue crack propagation. The main effect of the environment seems to be a reduction in the mechanical stresses needed to propagate the fatigue cracks. It is noticeable that the Leinenbach et al. [129] proposed a different approach to evalu- trajectory of the crack growth runs almost parallel to the Kmax axis ate the influence of corrosion on the fatigue behavior of titanium in the high crack growth regime for the alloys exposed to air and alloys. According to the authors, corrosion potential measurements saline solution. This behavior is typical of type V fatigue crack can give in situ information on surface damage during rotating growth [63], and is related to stress corrosion control; that is, crack bending in a corrosive environment. The formation of intrusions growth is governed rather by Kmax than DK⁄, meaning that the and extrusions during cyclic stresses lead to the creation of new stress level is more important than its frequency. non-passivated surfaces in metallic implants. Crack initiation and Even though the article by Lee et al. [128] is not focused directly propagation act in the same manner. When the material is in con- on biomedical applications, it exposes the crucial role that environ- tact with an aggressive environment, the passive film is damaged, ment and microstructure play on the fatigue crack growth behav- thus creating fresh new passive film-free surfaces. Consequently, ior of titanium alloys. Such an approach is hardly ever found in the the open circuit potential is shifted to more negative values. literature regarding biomedical metallic materials. In this respect, Leinenbach et al. [129] employed electrochemical noise to simulta- it is evident that the development of new titanium-based orthope- neously follow the potential and corrosion current variations of a dic implants should encompass a more detailed analysis of fatigue Ti–6Al–7Nb biomedical alloy under cyclic deformation. In this crack growth behavior. The use of the methodology developed by electrochemical technique, the working electrode (the specimen Vasudevan and Sadananda [63] is efficient at identifying the mech- under testing) is electrically connected to a counter electrode. anisms involved in fatigue crack growth at specific environments The counter electrode consists of the same material with the same for a given material and microstructure. surface finishing condition as the working electrode. Hence, the R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962 947

Niinomi et al. [132] investigated the fatigue crack growth behavior of Ti–5Al–2.5Fe alloy in air and in Ringer’s solution. They tested the material with a microstructure composed of equiaxed a grains and fine a precipitates (denoted as material B). The fatigue behavior was compared to that of a Ti–6Al–4V ELI alloy with a typical Widmanstätten microstructure (denoted as material F). Both microstructures are shown in Fig. 14. For both materials, da/dN was higher in Ringer’s solution than in air in the Paris law regime. The crack closure effect was smaller in Ringer’s solution than in air. The fracture surface would become smoother in Ringer’s solution as a consequence of corrosion. Thus, when a fatigue crack is nucleated, the fracture surface is more easily dissolved in physiological solution than in air. Based on the da/dN vs. DK plots, the corrosion-assisted fatigue crack growth of both materials may be classified as type II behavior (pure corrosion fatigue process) according to the classification scheme of Vasudevan and Sadananda [63].

Fig. 13. Corrosion potential (U) and corrosion current (Icorr) during fatigue crack growth of Ti–6Al–7Nb alloy (from Leinenbach et al. [129]; reprinted with 2.4.4. In vivo corrosion mechanisms and failure analysis permission from Elsevier). In addition to the study of the fatigue properties of biomedical alloys through laboratory testing, failure analysis of prematurely counter electrode will be in the same electrochemical state as the fractured implant materials is also of prime importance to the working electrode and no net current will flow between them understanding of the mechanisms that govern environment-as- while the surface of the loaded specimen is undamaged. Upon sisted fatigue failure of metallic implants under real conditions. starting the test, the local destruction of the passive film on the Currently, fracture of modern orthopedic implant devices is con- surface of the working electrode gives rise to variations in both sidered a rare complication in total hip arthroplasty [133]. How- the potential and the corrosion current. If oxygen is dissolved in ever, contemporary modular designs used in this type of surgical the electrolyte, a new oxide film may develop immediately after procedure allow micromotion at the Morse taper junctions during activation of the surface. As a result, the ion flux into the electro- mechanical loading [134]. Furthermore, crevices are formed be- lyte is reduced, the corrosion current drops and the potential in- tween the taper junctions, creating conditions for the penetration creases. A constant decrease in potential and a constant increase of body fluids that remain stagnant. As a consequence, crevice cor- in the corrosion current indicate that more new surfaces are cre- rosion may develop. As corrosion advances in this confined area, ated and subsequently repassivated within a certain period of time. oxygen concentration drops, leading to an excess of metallic ions. Thus, the evaluation of fatigue crack nucleation and growth can be Negatively charged chloride ions migrate from the physiological carried out through potential and current measurements during solution to the crevice. Hydrochloric acid then begins to form in- the number of stress cycles in a fatigue test. side the crevice, dissolving the protective oxide film on the surface An example of this approach used by Leinenbach et al. [129] of the implant. This type of corrosion has been found in stainless with their Ti–6Al–7Nb orthopedic alloy is shown in Fig. 13. First, steel, cobalt–chromium and titanium implants [135,136]. the crack grows along the surface of the alloy. The small current The combination of crevice corrosion and fretting fatigue is a peaks indicate that the crack growth rate is low. This would denote critical issue in the event of mechanical failure of metallic im- that only a reduced fraction of the surface of the titanium alloy is plants. A recent report by Paliwal et al. [137] confirmed the associ- activated during each stress cycle. As the crack becomes longer, ation of fretting fatigue and corrosion in the fracture of three the load at the crack tip increases and so does its growth rate. This uncemented Ti–6Al–4V alloy modular total hip stems. The general regime is evidenced by a sudden rise of the current peaks and a mechanism of failure begins with the formation of microcracks on sharp reduction in the corrosion potential. Final fracture leads to the surface of the implant during fretting corrosion. Microcracks a distinct peak of the corrosion current due to the relatively large then propagate because of fatigue until the load reaches the fatigue activated surface characteristic of the last stress cycle before strength of the material, leading to brittle fracture. Rodrigues et al. failure. The validity of this mechanism was confirmed through [138] observed the susceptibility of Ti–6Al–4V/Ti–6Al–4V modular microscopic investigations by scanning electron microscopy. This taper interfaces in the stem to fretting and crevice corrosion. More- approach was further developed by the same group in other over, hydrogen embrittlement was also found due to the typical publications [130,131]. mechanism of crevice corrosion in these devices. Yokoyama et al.

Fig. 14. Microstructures of B and F alloys (from Niinomi et al. [132]; reprinted with permission from Elsevier). 948 R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962

[139] have also shown that hydrogen absorption plays an impor- nitrogen alloying increases pitting corrosion resistance of austenit- tant role in the corrosion fatigue process of titanium dental ic stainless steels in neutral and acidic environments [155]. This implants. effect has been discussed by many authors [156–158]. It is virtually impossible to fully reproduce the complexity of Fatigue crack initiation has been correlated to the presence of in vivo applications of metallic implants in laboratory fatigue tests. pits in metallic materials [8,159]. Repassivation of mechanically Variable frequency of oscillating stresses, patient’s health and life- depassivated slip steps may not occur when the material is at a po- style, and accuracy of the surgical procedure are factors that pose tential in the metastable pitting range. Localized attack is then serious difficulties at reproducing real service conditions of an im- likely to occur at these active sites. This mechanism is known as plant through in vitro tests. Recent developments of titanium- the slip step dissolution model [160]. Pits are formed in the pro- based implants have attempted to give a step in this direction cess, acting as stress raisers and preferential sites for fatigue crack [140,112,141]. However, a complete evaluation of fatigue life of initiation. Thus, as molybdenum and nitrogen improve the pitting metallic implants has not been performed to date. Accurate simu- corrosion resistance of austenitic stainless steels, adding such ele- lation of the complex natural body fluid, composed of proteins, en- ments would also be an effective way of improving corrosion fati- zymes and saline species that act and interreact simultaneously, gue resistance of stainless steel implants. Additionally, as depicted and of the stresses that combine fretting and fatigue conditions by Begum et al. [142], the low stacking fault energy (SFE) of nitro- is still to be done, as noted by Fleck and Eifler [3]. There is a clear gen is closely related to its positive effect on the corrosion fatigue need to drive future developments of biomedical alloys along this resistance of austenitic stainless steels. Low SFE is associated with route. Moreover, the lack of studies on fatigue crack propagation of an increase in the amount of planar slip. This, in turn, increases the titanium alloys in simulated body fluids is surprising. There has resistance of the material to localized strain and hinders the forma- been no systematic investigation of the correlation between micro- tion of PSBs, thus hindering fatigue crack initiation [161]. structures, surface finishing and fatigue crack growth of biomedi- In spite of the widespread use of austenitic stainless steels as cal titanium alloys under real service conditions, i.e. simulating load-bearing devices in biomedical applications [162], there are crevice corrosion in the physiological medium composed of a mix- relatively few investigations on the corrosion fatigue phenomena ture of proteins, enzymes and saline species at 37 °C, and with dif- associated with fatigue life, especially the fatigue crack growth of ferent conditions of wear and frequency of the cyclic stresses. This these materials. One exception is the report by Giordani et al. lack of information is not exclusive of titanium alloys. The same [12]. They investigated the corrosion fatigue crack initiation of a deficiency is encountered in the literature concerning stainless high-nitrogen ISO 5832-9 austenitic stainless steel. This type of steels, cobalt–chromium and biodegradable magnesium alloys, as steel was developed to replace the conventional ASTM F-138 grade, described in the next sections. owing to its higher pitting corrosion resistance [163]. Their fatigue data were expressed as S–N curves obtained in air and 0.9 wt.% 2.5. Stainless steels NaCl solution at 37 °C. The saline environment reduced the fatigue life of the specimens. The mechanism of failure was not attributed The stability of the passive films formed on the surface of stain- to the primary onset of pitting corrosion. No evidence of pitting less steels strongly affects the corrosion fatigue behavior of these was observed on the surface of the specimens tested in the aque- materials [142]. Thus, the chemical composition is decisive in ous environment. The preferential sites for crack initiation were determining fatigue properties. In this context, it has been reported coarse precipitates, composed mainly of Nb and Cr (a complex ni- that the addition of molybdenum and nitrogen to austenitic stain- tride of Nb and Cr called the Z-phase), and Al-rich non-metallic less steels is an effective method for improving their corrosion fa- inclusions, as shown in Fig. 15. This mechanism operates for both tigue resistance. Analytical studies by means of Auger electron fatigue and corrosion fatigue failure. According to the authors, spectroscopy and X-ray photoelectron spectroscopy show that the first cycles of stress plastic deformation lead to the fracture the oxide layer on austenitic stainless steels has a duplex structure of many coarse precipitates, forming geometric discontinuities at [143,144]. It is known that the composition of the passive film is these sites. The electrolyte then penetrates through such disconti- strongly influenced by the composition of the alloy and the pH of nuities, reaching the bare metal underlying the fractured passive the solution [145,146]. It is also well documented that the passive film. As a result, a crevice corrosion mechanism is developed, layer on these materials is composed of an inner chromium-rich region of a few atomic layers in contact with the metallic substrate and an external iron-rich layer at the interface between the film and the electrolyte [147,148]. Molybdenum is known to have beneficial effects on the pitting corrosion resistance of austenitic stainless steels [149]. The high energy of Mo–Mo bonds increases the activation energy barrier for dissolution, thus enhancing resistance to pit initiation. Nitrogen influences passivity of austenitic stainless steels [150,151], increasing the stability of the passive layer. This effect has been explained through several mechanisms. Jargelius-Petersson [152] found that the primary role of nitrogen is in pit repassivation rather than in the onset of pitting corrosion. A synergism between molybdenum and nitrogen was also proposed based on the fact that the combination of these elements was effective at eliminating current transients during passive current density measurements in molybdenum-bearing stainless steels. Other mechanisms are based on the formation of ammonium ions or nitrate/nitrite ions [142,153] and the enrichment of nitrogen on active surfaces [154]. The specific mechanism through which nitrogen acts may Fig. 15. Coarse Z-phase precipitate and Al-rich non-metallic inclusion acting as depend on the characteristics of the electrolyte in contact with fatigue crack initiation sites in ISO 5832-9 stainless steel (from Giordani et al. [12]; the stainless steel such as pH. However, it is well recognized that reprinted with permission from Elsevier). R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962 949 continuously impeding repassivation inside the discontinuities. Tavares et al. [166] investigated the causes of failure of stainless Consequently, anodic dissolution of the metallic material proceeds steel orthopedic implants (dynamic compression plates) removed with the evolution of the cyclic loading, giving rise to an autocata- from patients after premature fracture. Failure was related to a lytic process. Therefore, corrosion accelerates fatigue failure by combined corrosion fatigue process accelerated by the poor surface intensifying the concentrations of stress and strain at the precipi- finishing of the prostheses that acted as stress concentration sites tates and non-metallic inclusions. to the nucleation of fatigue cracks. Crevice corrosion in the bolt/ The corrosion fatigue failure of an orthopedic stainless steel im- plate contact also contributed to the corrosion-assisted fatigue fail- plant was studied by Amel-Farzad et al. [164]. The device was a ure. This article reveals the decisive role that surface finishing DCS barrel plate, which had been implanted in the patient’s thigh plays on the corrosion fatigue behavior of metallic implants. Even for two years. A top view of the fractured plate is shown in when surface corrosion is not present, fatigue failure of stainless Fig. 16. Pits and cracks are visible on the surface of the implant steel compression plates may occur as a result of cleavage decoher- with unaided eye. Failure was found to be due to corrosion fatigue ence, as reported by Triantafyllidis et al. [167]. assisted by a crevice corrosion mechanism. The material used to Xie et al. [168] extended the analysis of the mechanisms that manufacture the plate had a chemical composition that was totally govern corrosion fatigue of stainless steel implants. They con- inadequate for such biomedical applications, with much lower cluded that the initiation of corrosion fatigue cracks is due to inter- nickel, chromium and molybdenum contents than those recom- granular corrosion inside pits. The initial cracks are generated mended by the ASTM F-138 standard. The low quality of the mate- along attacked grain boundaries and propagate intergranularly. rial led to the onset of pitting corrosion inside the crevices between In a next step, transgranular cracks are formed from the initial the screws and the plate. Then cyclic loadings nucleated cracks at intergranular ones due to the local stress field at the crack tip. the localized corrosion sites. Cracks propagated with the increase One of the transgranular cracks dominates fatigue crack propaga- in stress cycles and, when a critical size was reached, the final tion, leading to the final failure. Dislocation pile-up is also closely fracture occurred. related to this mechanism. The authors observed that dislocations Swarts et al. [165] carried out a retrieval study of fractured high- multiply and pile up along grain boundaries during corrosion fati- nitrogen stainless steel tapered hip stems. The authors emphasize gue tests of surgical stainless steels in laboratory. The longer the the superior corrosion resistance of this grade to conventional cyclic stress, the more pronounced is this effect. As a result, the 316L surgical materials and point that the occurrence of failure is chemical activity of stainless steels will increase along the high dis- rarely reported for these devices. However, six different cases of fail- location density grain boundaries. Intergranular corrosion is facili- ure of tapered hip stems manufactured from this material had been tated at these sites and propagates rapidly to become a dominant forwarded to the authors’ laboratory for an accurate evaluation of transgranular crack under cyclic stresses. the main causes involved in the fractures. The failure mechanism Bolton and Redington [169] conducted an investigation of the was identified as corrosion-initiated fatigue for all stems. The fatigue crack growth behavior of two surgical grade (316L and authors conducted potentiodynamic polarization measurements 316LVM) stainless steels in physiological solution. This is a rare to assess the pitting corrosion resistance of the materials. The re- example of the use of the FCP approach to the evaluation of the fa- sults pointed to a high resistance to the onset of pitting. The pres- tigue behavior of surgical stainless steels. For both materials, crack ence of localized corrosion attack at the surface of the stems was growth rates increased in the presence of the saline solution in thus considered surprising by the authors. The combined effect of comparison with air at low stress intensities. This trend was re- wear and crevice corrosion was hypothesized as a possible cause versed at high stress intensities. The da/dN vs. DK plots for 316L for the localized corrosion signs on the stems: fatigue cracks had ini- stainless steel in air and in Ringer’s solution is shown in Fig. 17. tiated at these sites, leading to catastrophic failure. The crack growth behavior of the materials may be classified as A further explanation for the failure mechanism was given type II, i.e. a true corrosion fatigue, according to the classification based on the microstructure of the stems, especially the grain size. scheme of Vasudevan and Sadananda [63]. The decrease in da/dN The microstructure of the stems was found to present an average with increasing DK may be related to the secondary cracking that ASTM grain size in the range specified by the ISO 5832-9 standard occurred along striations. The authors found that above DK = (the maximum allowable grain size is ASTM 4, based on the mea- 30 MPa.m1/2 a fully developed striation appearance formed across surements described in ASTM E-112 standard). However, the grain the fracture surface and the striation spacing increased with stress size distribution was heterogeneous, and some grains at the sur- intensity. Consequently, the opening up of a fatigue crack along the face of the stems were coarser than the maximum size specified base of striations suggests that changes took place during the re- by the standard. The authors suggest that grain size heterogeneity versed plastic flow. As a result, the shape of the fatigue cracks has the potential to reduce the fatigue strength of the alloys, was altered, as was the morphology of the striations, leading to a increasing the probability of fatigue failure. A careful control of decrease in the crack growth rate. This analysis was supported the grain size through the combination of specific mechanical by SEM images of the fractured surfaces. processing (cold work) and heat treatment (recrystallization) oper- The number of scientific investigations of the fatigue behavior ations is suggested to avoid this problem. The analyses by Swarts of stainless steel implants is very low in comparison with those et al. [165] highlight the importance of controlling specific of titanium-based alloys, especially during the last decade. This is microstructural aspects of austenitic stainless steel in order to at- probably related to the well-documented limitations of stainless tain optimized corrosion fatigue properties. steels owing to corrosion resistance in body fluids, the release of potentially damaging ions to adjacent tissues and the stress shielding effect [170–173]. Titanium-based biomaterials are known to behave better in all these aspects [48,174]. However, permanent prostheses made of stainless steel are still widely used due to cost savings [175,176]. This scenario is likely to continue for an indefi- nite period. While the study of fatigue crack growth behavior of the stainless steels used in nuclear power or chemical plants has received considerable attention [177–179], a systematic investiga- Fig. 16. Top view of the stainless steel fractured plate (from Amel-Farzad et al. tion of this subject regarding biomedical applications has been [164]; reprinted with permission from Elsevier). neglected. 950 R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962

showed that pitting corrosion was present at the fracture surface as well as fatigue striations. It is not only the design of Co–Cr–Mo implant devices inﬂuences the fatigue behavior of these components; their microstructure is also of prime importance. A typical as-cast microstructure of cobalt-based surgical alloys presents a dendritic a-fcc metastable

matrix and precipitates composed mainly of blocky M23C6 carbides, which appear in interdendritic regions and grain boundaries [190]. A lamellar phase has also been observed in the grain bound-

aries. Furthermore, an intermetallic sigma (r) phase and M6C carbides may be present as minor constituents of the secondary phase aggregates [191]. An example is shown in Fig. 18. Carbide precipitation accounts for the main strengthening mechanism of the alloys in the as-cast condition. Coarse blocky carbides act as sources of dislocations and stacking faults when the material is subjected to mechanical loads. As suggested by Dobbs and Robertson [192], the lamellar constituent is the most detrimental phase of the microstructure. When it is removed by heat treatment, the tensile strength, ductility and fatigue life are increased. Hence, it is possible to improve the fatigue behavior of cobalt-based implants through the careful control of alloy composition and heat treatment operations. Alloy composition is primarily related to the carbon content in the alloy, which provides the basis for carbide formation during cooling from the melt. However, other elements can improve the overall mechanical properties of Co–Cr–Mo biomedical alloys and, consequently, their fatigue properties. Zhuang and Langer [193] observed that nickel additions

Fig. 17. Fatigue crack growth rate as a function of stress intensity range for 316L stainless steel in air and in Ringer’s solution (from Bolton and Redington [169]; reprinted with permission from Elsevier).

2.6. Co–Cr–Mo alloys

Co–Cr–Mo alloys are known to present high wear resistance and good mechanical properties under static loadings [180]. In addition to these features, intrinsic biocompatibility and excellent corrosion resistance are also key factors that determine the use of these alloys for orthopedic applications [181]. Moreover, the fatigue strength of cobalt-based biomaterials exceeds that of titanium alloys and austenitic stainless steels, depending on the manufacturing methods employed to produce the alloys [182,183]. Nevertheless, fatigue failures of Co–Cr–Mo orthopedic devices have been reported [184–186]. Corrosion processes are often associated with these catastrophic failures. Gilbert et al. [187] analyzed the fracture causes of two modular total hip prostheses, consisting of a wrought cobalt alloy (ASTM F799) head coupled to a wrought cobalt alloy stem. The two fractures occurred in the neck region, outside of the taper junction between the head and neck components. Modularity is known to favor the onset of crevice corrosion in orthopedic metallic implants [188]. In fact, Collier et al. observed several pits on the fracture surface of the prostheses which were created inside the crevices between the head and the stem. Manufacturing defects such as porosity and segregation may have also contributed to the corrosion-assisted fatigue failure of the prostheses. Sudhakar [189] reported on the corrosion fatigue failure of a Co–Cr–Mo alloy (Vitallium 2000) used as a bone plate. The fracture was found to occur at the interface between the screws and the Fig. 18. Microstructure of a Co–Cr–Mo surgical alloy (ASTM F75) (from Ramírez- plate, a typical site for the onset of crevice corrosion. Sudhakar Vidaurri et al. [190]; reprinted with permission from Elsevier). R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962 951 improved the fatigue crack growth resistance of a Co–Cr–Mo bio- sion resistance in physiological fluids have been evaluated [217– medical alloy by increasing its stacking fault energy. This prevents 219]. the reduction of ductility of the alloy [194] and is beneficial to the The corrosion fatigue of structural magnesium alloys has been workability during subsequent cold working operations, such as studied by several authors in NaCl and borate-buffered solutions. forging [195]. Heat treatments, in turn, provide specific micro- These investigations have generally focused on applications of structural alterations that are directly related to the mechanical magnesium alloys in the electronic, automotive and aerospace behavior of the material under both static and cyclic stresses [196]. industries [220–223], and have used the cumulative fatigue dam- Despite the well-documented failures of cobalt-based implants age (CFD) approach to assess the corrosion fatigue behavior of due to a corrosion fatigue mechanism, few attempts have been magnesium alloys. However, it is known that crack initiation life conducted to predict the fatigue crack growth of these materials is significantly reduced due to pit formation in a corrosive environ- under the influence of a corrosive environment or even in air ment [224]. The fatigue strength of magnesium alloys is signifi- [197–199]. Marrey et al. [200] conducted an investigation into this cantly reduced in humid environments, and the fatigue limit subject. They developed a damage-tolerant analysis for the fatigue drops drastically in NaCl solution. This reduction is caused by the crack growth behavior of a Co–Cr–Mo alloy used as a cardiovascu- formation of pits on the surface of the alloy, the initiation of fatigue lar stent. The mathematical model was compared with experimen- cracks at these sites and their subsequent growth [225], leading to tal results obtained in Ringer’s solution at 37 °C. The authors final catastrophic fracture. Thus, the CFD approach, which focuses predicted that, provided that all flaws greater than 90 lm were de- on fatigue strength and fatigue life rather than on crack growth, tected prior to the implantation of the stent, in vivo fatigue failure is not suitable to fully understand the corrosion fatigue behavior was unlikely to occur. Niinomi [1] reported such a reduction in fa- of magnesium alloys. tigue strength of Co–Cr–Mo alloys in physiological solution in com- Despite this, an alternative approach – the fatigue crack propa- parison with fatigue tests conducted in air. However, Niinomi’s gation (FCP) approach – has been used by only a few authors report is based on the CFD approach, ignoring the fatigue crack [223,224]. Rozali et al. [224] studied the fatigue crack growth of growth behavior of the metallic implants. If one considers the AZ61 alloy in NaCl 3.5 wt.% solution at different frequencies. They widespread use of cobalt-based orthopedic devices, it is clear that observed that the fatigue crack growth rate was higher when the research on the corrosion fatigue behavior of these materials needs alloy was immersed in the NaCl solution in comparison with a to evolve. In particular, the fatigue crack propagation approach low humidity environment. Moreover, fatigue cracks were shown simulating crevice corrosion conditions in physiological solutions to propagate under lower load amplitudes in the saline environ- has been hardly considered in the literature. ment. Nan et al. [223] proposed a mathematical law to model the corrosion-assisted fatigue crack growth of AZ31 alloy in 3 wt.% NaCl solution. Experimental da/dN vs. DK plots and modeled 2.7. Magnesium-based alloys data corresponded well. Despite the relatively high number of scientific reports on the Magnesium alloys are currently considered for applications as corrosion behavior of biomedical magnesium-based alloys, load-bearing implant devices such as plates, screws and pins for corrosion fatigue has received little attention. Although these repairing bone fracture [201]. This trend arises from the low corro- materials are designed to dissolve within the human body in sion resistance of magnesium alloys in aqueous environments accordance with their application as temporary fixation compo- [202]. While for most engineering applications the susceptibility nents, they have to maintain suitable mechanical strength during to corrosion is a critical limitation of these materials, for biomedi- the healing of the fractured bone [226]. Hence, corrosion fatigue cal purposes it is a desirable property. If the material is employed should be thoroughly evaluated to ensure that the orthopedic de- as a fixture device, degradation may be beneficial to the patient, vice will perform suitably and safely. Recently, Gu et al. [227] since the device will be absorbed by the body and hence the need provided a proper analysis of this problem and took a step to- for a new surgical procedure to remove the device will be avoided wards a deeper understanding of this phenomenon by evaluating [203]. the corrosion fatigue behavior of AZ91D and WE43 biomedical From this standpoint, biodegradable magnesium-based im- magnesium alloys. They compared the fatigue properties of both plants are very attractive since the in vivo corrosion of these mate- materials in air and in simulated body fluid (SBF) at 37 °C rials generates mainly soluble, non-toxic products. This has been through S–N curves. The fatigue limit was found to significantly well studied in the literature [204–206]. However, hydrogen gas decrease for both alloys in SBF solution, as shown in Fig. 19 for also evolves during in vivo degradation of Mg alloys. If the degra- the AZ91D alloy. The mechanism of fatigue failure was associated dation is too fast, hydrogen may accumulate as subcutaneous gas with the stress concentrations in casting defects such as internal bubbles. Moreover, the implant device may lose its mechanical pores, and with crack initiation and growth at surface pits integrity before the effective healing of the fractured bone formed in the SBF solution. Chamos et al. [228] also showed [207,208]. Hence, controlling the corrosion rate of magnesium- the major role played by pitting corrosion in the fatigue failure based biomedical alloys is of utmost importance. of AZ31 magnesium alloy. Several studies have concentrated on this subject, investigating The report by Gu et al. [227] clearly shows the harmful effect the effect of different alloying elements and coatings on the corro- physiological solutions can have on the fatigue behavior of magne- sion properties of biodegradable magnesium alloys [209–212]. Re- sium alloys. The importance of low cycle fatigue is also outlined, ports have focused mainly on aluminum- and rare earth (RE)- since final failure may occur within the typical number of cycles containing magnesium alloys due to the overall positive effect that that a temporary implant may be expected to withstand during these elements have on the corrosion resistance of magnesium the healing of the fractured bone. In this scenario, in addition to [213]. However, Witte et al. [201] have recently recommended that the evaluation of the corrosion fatigue behavior by the CFD ap- MgAl and MgRE systems should not be used for biomedical appli- proach, it is very important to consider fatigue crack propagation, cations due to indications of possible harmful effects of these ele- analyzing the da/dN vs. DK plots with regard to the growth of fati- ments to osteoblasts and the onset of Alzheimer’s disease or gue cracks under simulated physiological conditions. This ap- cytotoxicity to specific cell lines [214–216]. Consequently, new proach is not found in the literature for the modern Al-free Al- and RE-free biodegradable magnesium alloys have recently biodegradable magnesium alloys or even for the conventional Al- been developed for biomedical purposes and their in vitro corro- containing materials such as AZ91D, AZ31 or AZ61. 952 R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962

one considers metallic implants, the lack of information is still more pronounced. Therefore, this section deals with the analysis of prevention and mitigation methods used to improve the fatigue properties of biomedical alloys. The more conventional methods, based on compressive residual stresses and hard coatings, are depicted in the ﬁrst two subsections, then the potentialities of surface nanocrystallization are discussed. As will become evident the reader, there is an undeniable scarceness of corrosion fatigue studies even for the conventional mitigation methods. This scarceness is still more signiﬁcant for alloys with a nanocrystallized surface.

3.1. Compressive residual stresses

Compressive residual stresses are created by finishing processes that impart localized plastic deformation to the material surface. At the surface, deformed regions tend to expand, but the un- strained regions more distant from the surface restrain this expan- sion [245]. Thus, compressive strains are generated at the deformed surface. The compressed surface is effective at prevent- Fig. 19. S–N curves for die-cast AZ91D at room temperature in air and in SBF at 37 °C (from Gu et al. [227]; reprinted with permission from Elsevier). ing fatigue failure since the fatigue cracks are generally initiated at zones subjected to tensile stresses at the free surface of a metallic material [237]. From the foregoing context, it is clear that research on the cor- The most widespread method to obtain compressive residual rosion fatigue behavior of magnesium-based biomaterials is still stresses is shot peening. This is based on the bombardment of a incipient. The interaction between pitting corrosion and fatigue metallic surface with small glass, ceramic or metallic spherical par- failure is clear and relatively well known from investigations in ticles (shots), which undergo multiple and repeated impacts (peen- NaCl solution. However, fatigue crack propagation behavior needs ing) against the surface of ther alloy [246]. As a consequence, the to be studied for a wide variety of biodegradable Mg alloys. Even dislocation density is increased in the near-surface zones, as are reports on the CFD approach are not found for Al-free alloys. Fur- roughness and hardness [247]. Surface roughening is considered thermore, the relationship between microstructure and corrosion to be deleterious to the fatigue strength of shot-peened materials fatigue behavior in physiological medium has been ignored in [248]. Hence, shot peening has both beneficial and harmful effects the literature. Several aspects, such as grain size, cold work, twin- on the fatigue properties of metallic alloys. The optimization of ning, crystallographic texture and the content of specific alloying parameters such as peening intensity, peening coverage, satura- elements such as Ca, Zn and Mn, have been evaluated separately, tion, shot material, shot size, shot velocity, shot hardness and time i.e. focusing on their individual effects on corrosion or fatigue of peening is essential to reach the best conditions of compressive behavior [203,229–231], whilst their combined action has been residual stresses and surface roughening [249]. The determination ignored. This scenario should be modified by combining FCP and of the optimum parameters is a time-consuming task. Laser shock CFD approaches through the investigation of the relationship peening (LSP) has been developed as an alternative finishing pro- between the microstructure and corrosion fatigue mechanisms of cess in order to achieve deeper compressive residual stresses and biodegradable magnesium alloys. This is an imperative step avoid the deleterious effect of exaggerated surface roughening towards the development of fail-safe orthopedic components. after shot peening [250]. LSP imparts plastic deformation to the treated surface through the formation of a plasma cloud that is generated by a laser pulse and travels into the workpiece as a 3. Mitigation shock wave, creating compressive residual stresses [251]. Both shot peening and LSP have been used to increase the fatigue According to Bayraktar et al. [232], the microstructures of alloys strength of metallic implants, LSP being investigated more recently and their manufacturing processes are closely related to fatigue [252–255]. Semlitsch et al. [252] assessed the fatigue behavior of failure. Prevention of fatigue failure starts through the careful con- Ti–6Al–7Nb hip stems through S–N curves. The endurance limit of trol of alloy composition and processing. Microstructure-related shot-peened components was increased in comparison with influences have been addressed in the previous section. Surface ef- polished oxygen-diffusion-hardened specimens. Furthermore, the fects also determine the fatigue behavior of metallic materials fatigue strength was found to be sensitive to surface roughness, [233]. It is widely recognized that surface finishing operations with a reduction associated with the rougher surface. may be favorably used to increase the fatigue strength of metallic Papakyriacou et al. [123] compared the corrosion fatigue behav- materials [234–236]. Conventional methods for fatigue life ior of Ti–6Al–7Nb and commercially pure Ti specimens with two improvement are based on the introduction of surface compressive different surface conditions: ground and shot peened. The fatigue residual stresses and the deposition of hard thin films [237–240]. tests were conducted in a physiological solution that simulated Metallic implants have benefited from both of these methods. A inflammatory conditions at the oral cavity. The shot-peened spec- hot topic has emerged in the last decade: surface nanocrystalliza- imens were less sensitive to fatigue fracture in the corrosive solution as an alternative way to increase the corrosion resistance of tion than ground materials. metallic alloys. This approach has proved to be efficient in a variety Azar et al. [253] investigated the effect of the time of peening on of materials, such as stainless steels, zirconium, titanium alloys the corrosion and fatigue behavior of a surgical AISI 316L stainless and nickel alloys [241–244], regarding their applications as bioma- steel. These properties were evaluated separately, i.e. some speci- terials or not. mens were subjected to fatigue tests by a rotating, bending method Despite the large number of investigations on the effect of in laboratory air and others underwent the electrochemical mea- nanocrystallization on the corrosion resistance of specific metallic surements in Ringer’s solution at 37 °C. Hence, the study was not materials, corrosion fatigue has hardly been evaluated at all; and if focused on the assessment of the combined action of cyclic loading R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962 953 and corrosion attack on the shot-peened material. Nevertheless, it provides information useful in elucidating the influence of the shot-peening process on the performance of surgical stainless steel regarding its corrosion and fatigue behavior. The number of cycles to failure increased with the time of peening, as observed in the fatigue tests. This result was due to the high compressive residual stresses introduced for longer times of peening, postponing the initiation of fatigue cracks. Electrochemical measurements pointed to an increase in the corrosion rate and a reduction in the breakdown potential after up to 15 min of shot peening. The increase in surface roughness associated with the plastic deformation caused by the impact of the peening medium was closely related to the reduction in corrosion resistance observed with up to 15 min of peening. However, the surface roughness was found to decrease for longer times of peening, and the compressive residual stresses increased. These two effects have been correlated with a reduction in the corrosion rate in austenitic stainless steels [254,255]. This approach reveals the strong effect of process parameters on the performance of shot-peened biomedical alloys. Nevertheless, despite the usefulness of these findings, the actual material response could Fig. 20. S–N curves of commercially pure grade 4 Ti with different surface finishing (from Pazos et al. [256]; reprinted with permission from Elsevier). only be perceived if the simultaneous effect of fatigue and corrosion had been investigated by the authors. Wieser [254] made a valuable contribution to this topic by studying the effect of conventional shot peening on the corrosion fatigue of the austenitic decrease in fatigue strength. This was caused by the introduction of X2CrNi–Mo-18-15-3 surgical steel. Wieser conducted fatigue tests several surface defects, mainly represented by microholes and in laboratory air and in 0.9 wt.% NaCl solution at 37 °C with shot- intergranular corrosion, as observed through SEM micrographs. peened specimens and electrolytic polished specimens. The effect These defects acted as stress raisers, giving rise to the early initia- of the combined surface treatments (shot peening followed by tion of fatigue cracks. The treatments that involved plastic defor- electrolytic polishing) on the fatigue behavior of the specimens mation, i.e. shot blasting and the dual treatment, introduced was also evaluated. Without the influence of corrosion, the shot- compressive residual stresses in the surface of the metallic mate- peened specimens presented higher fatigue strength than the elec- rial. Consequently, fatigue strength was higher in comparison with trolyte-polished ones. By combining the two finishing processes, the etched material. However, the as-machined material was more the fatigue strength was further improved. The results obtained resistant to fatigue than the shot-blasted one. Shot blasting created in the physiological solution evidenced a pernicious effect of shot notch-shaped surface defects that were preferential sites (stress peening under corrosion fatigue conditions. The fatigue strength raisers) for the nucleation of fatigue cracks. This effect counteracts of the shot-peened specimens was reduced by 30% in the saline the benefit of the compressive residual stresses generated during solution, while for the electrolyte-polished specimens the reduc- the treatment. The combination of shot blasting and acid etching tion was only 13%. According to the author, after shot peening, produced the best performance amongst the three surface treat- the surface roughness increases as well as the specific surface area ments. This result resembles the findings by Wieser [254],who of the material. As a result, localized corrosion attacks occur at the showed that a dual treatment consisting of shot peening and elec- narrow depressed sites in the rough surface, making the material trolytic polishing increased the corrosion fatigue resistance of a more susceptible to the initiation of fatigue cracks. Moreover, the surgical stainless steel. dislocation density increases due to the strain hardening effect of The increase in fatigue resistance reported by Pazos et al. [256] the shot-peening process, shifting the corrosion potential toward was only marginal in comparison with the reference as-machined the anodic direction and further decreasing the corrosion resis- finishing. In addition, it is important to emphasize that the results tance of the peened surface. The specimens subjected to the com- were obtained in air. Thus, it is not unlikely that, under a combined bined action of shot peening and electrolytic polishing presented corrosion fatigue mechanism, the performance of the biomedical the best behavior under the corrosion fatigue conditions. The fati- material would be further reduced. It is notable, then, that surface gue strength was decreased by only 4% in comparison with the re- treatments have to be carefully evaluated regarding their influence sult obtained in air. Surface smoothing due to electrolytic polishing on the fatigue behavior of metallic implants and, most importantly, was very important in preventing the harmful effects of corrosion- corrosion processes should always be considered in order to assisted fatigue crack initiation. achieve reliable results. Pazos et al. [256] evaluated the fatigue strength of annealed Shot peening has also been performed on magnesium-based al- commercially pure grade 4 Ti subject to acid etching (9 M sulfuric loys. Most reports are concerned with aerospace or automotive acid solution at 60 °C for 15 min), shot blasting with Al2O3 particles engineering applications [257–259], while biomedical use has re- and a dual treatment consisting of the shot blasting treatment fol- ceived only minor attention [260]. Irrespective of the application, lowed by the acid-etching treatment. The as-machined surface was it is important to concentrate on the outputs of the shot-peening used as a reference condition. The results are summarized in process regarding its effect on the fatigue properties of magnesium Fig. 20. The tests were conducted in air at room temperature. alloys. In this respect, Zhang et al. [257,258] evaluated the effect of Hence, only the fatigue behavior of the biomedical material was as- shot peening on the fatigue performance of a high-strength sessed, without any information on the corrosion-assisted fatigue wrought Mg–Al–Zn–Mn alloy (AZ80). The authors used the CFD ap- crack initiation. Despite the absence of investigations regarding proach and the tests were conducted in air. The results indicated corrosion-related phenomena, the results nevertheless showed that the fatigue strength and fatigue life of the AZ80 alloy could that surface finishing operations can be detrimental to the fatigue be improved by the compressive residual stresses introduced by properties of metallic implants. From the S–N curves shown in shot peening. The extent to which the fatigue performance can Fig. 20 it is clear that the acid-etching treatment led to a significant be improved depends on the peening medium and is directly 954 R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962 related to the degree of compressive stresses generated during the nitriding have been reported by several authors [261–263]. How- process. ever, perhaps the most versatile surface modification method is The influence of corrosion on the fatigue strength of the shot- based on the deposition of protective coatings on metallic im- peened magnesium alloy was not assessed by Zhang et al. plants, and involves the improvement of a variety of material prop- [257,258]. However, Khan et al. [259] provided a contribution to erties, such as biocompatibility, fatigue strength, corrosion and this field by investigating the corrosion fatigue behavior of a die- wear resistance. Dealing with all the possible deposition processes cast Mg–Al–Si–Cu–Mn–Zn alloy (AM60) subjected to shot blasting. and types of materials that may be employed as protective films for They showed that shot blasting was effective at improving the fa- biomedical alloys is not within the scope of this section. Instead, tigue performance of the alloy. S–N curves indicated that the fati- we aim to discuss representative data of selected examples, includ- gue strength was increased for the shot-blasted specimens in ing titanium alloys, austenitic stainless steels, Co–Cr–Mo and bio- comparison with the die-cast alloy in an NaCl solution. Fatigue degradable magnesium alloys. cracks originated from corroded areas at the surface of the mate- Even though the fatigue properties of metallic implants can be rial. The authors gave further insight into the fatigue behavior of greatly altered by a coating layer, it is not uncommon that the the AM60 by evaluating the fatigue crack growth rate of the die- primary goal of designing a coated orthopedic component is to cast material through da/dN vs. DK curves. The threshold stress increase its osteointegration capability rather than its fatigue intensity range (DKth) was found to decrease when the material strength [264,265]. In the same way, corrosion and wear resis- was in an NaCl environment (5 wt.% NaCl solution) in comparison tances are frequently focused on because of the direct relationship with low aggressive humid environments. The humid environ- between these properties and the biocompatibility of the biomed- ments consisted of a low relative humidity (55%) air atmosphere ical device [266–269], while the evaluation of the fatigue behavior and a high relative humidity (80%) air atmosphere. These results of the coated material is disregarded. However, several authors are reproduced in Fig. 21. It is concluded, then, that the fatigue have recognized the importance of considering the response of crack propagation rate is faster in a saline solution. This behavior coated metallic implants to cyclic loadings [270–272]. For exam- may be classified as pure corrosion fatigue (type II), according to ple, Wang et al. [270] studied the effect of titanium nitride (TiN) the classification scheme of Vasudevan and Sadananda [63]. Unfor- and diamond-like carbon (DLC) coatings on the fatigue, wear and tunately, this approach was not shown for the shot-blasted corrosion properties of AISI 316L surgical stainless steel specimens. specimens. The TiN layer was deposited using electron-beam plasma-assisted physical vapor deposition while the DLC film was produced by a 3.2. Coatings plasma-assisted chemical vapor deposition (PACVD) method. Fati- gue tests were performed with a ball-on-plate impact tester, as In addition to the introduction of compressive residual stresses schematically shown in Fig. 22. The test does not consider the at the surface of metallic implants by using shot peening and laser influence of corrosion on the cyclic loads applied to the coated bio- shock peening, other surface modification methods have been medical alloy; however, it shows the superior behavior of DLC- extensively used to improve the fatigue performance of these com- coated specimens in comparison with TiN-coated ones. Corrosion ponents. Passivation, electropolishing, chemical oxidation and tests were performed separately in a simulated body fluid. Both coatings significantly improved the corrosion resistance of the substrate. The increase in fatigue strength of metallic alloys due to the presence of PVD films has been reported by other authors [46,273–276]. Recently, Puchi-Cabrera et al. [277] showed that a DLC coating deposited by magnetron sputter ion plating increased the fatigue strength of 316L stainless steel in air and in 3 wt.% NaCl

Fig. 22. Representation of the impact tester used to assess the fatigue response of Fig. 21. Fatigue crack growth curves for as-cast AM60 alloy under different TiN and DLC-coated surgical stainless steel (from Wang et al. [270]; reprinted with environments (from Khan et al. [259]; reprinted with permission from Elsevier). permission from Elsevier). R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962 955 solution. Most importantly, the improvement in fatigue properties fatigue strength of the magnesium alloy in distilled water, whereas was more pronounced in the corrosive environment, as shown in a12lm multilayer DLC film enhanced it to nearly the same level Fig. 23. According to the authors, the benefit of PVD layers arise as in air. It is thus speculated that the thickening of multilayer due to the combination of high mechanical strength, compressive DLC films could completely suppress corrosion fatigue in aqueous residual stresses in the coating and good adhesion to the substrate. environments. This finding may be applied to enhance the corro- Other reports have demonstrated that PVD films enhance the wear sion fatigue properties of biomedical metallic alloys in physiologi- and fatigue strength of Co–Cr–Mo and Ti-based biomedical alloys cal medium. However, this approach is not found in the literature. [278–280]. However, if the coating delaminates because of poor Moreover, it is of fundamental importance in evaluating the fatigue adhesion to the substrate, then the presence of the PVD film may crack growth behavior of coated implants, yet very few such stud- reduce the fatigue strength of the coated metal instead of improv- ies are reported in the literature. ing it [281]. In an effort to enhance the wear and corrosion resistances of The results of Puchi-Cabrera et al. [277] show the good perfor- orthopedic devices such as femoral heads, alternative techniques mance that a thin, hard coating, such as a PVD film, can achieve to the PVD and CVD methods have been developed, focusing against corrosion fatigue in metallic alloys. Nevertheless, it is mainly on achieving better surface coverage (non-line-of-sight important to bear in mind that such layers have intrinsic limita- processes) and lower processing temperatures. In this regard, tions regarding their application as protective films on load-bear- sol–gel, electrolytic deposition and plasma immersion ion implan- ing implants. These include the formation of wear debris and the tation have been used to deposit ceramic-based hard films on bio- intrinsic brittleness of PVD thin films [282,283]. An overview of medical metallic alloys [285–288]. Although some successful such drawbacks and how to overcome them is given by Antunes applications are described in the literature regarding the improve- and De Oliveira [7]. ment of wear and corrosion resistances imparted by films depos- The suitability of using plasma-based technology of thin film ited by these techniques, the corrosion fatigue behavior is often deposition to enhance the fatigue properties of a magnesium- ignored, as occurs for the PVD- and CVD-coated metallic implants. based alloy was investigated by Uematsu et al. [284]. This work focuses on the application of AZ80A alloy as a structural material in 3.3. Surface nanocrystallization the aerospace and automotive industries. Biodegradable magnesium alloys are not intended for application in highly wearable The mechanical strength of metallic alloys can be controlled parts of an orthopedic device, such as a femoral head, nevertheless, through grain refinement. The conventional way to do this is to the results are important in that they evaluate the corrosion fati- combine plastic deformation and heat treatment. This is limited gue behavior of a CVD-coated metallic material. Few reports pro- to the microcrystalline scale, however, and the dependence of vide information on this subject. Thus, even though the work by the static mechanical strength on the grain size is given by the Uematsu et al. [284] is not directly related to the biomedical area, well-known Hall–Petch relationship [289,290]. Typical average it provides a valuable contribution to the understanding of the grain sizes are larger than 1 lm. The fatigue limit is often increased mechanisms involved in the corrosion fatigue of CVD-coated alloys. as well as the tensile yield strength after grain refinement at the Uematsu et al. evaluated the corrosion fatigue behavior of a microcrystalline scale [291]. The utmost barrier to this approach multilayer DLC film deposited on the AZ80A alloy (of which Al, is related to the development of nanocrystalline metallic materials. Zn and Mn are the main alloying elements) by a PACVD technique. The basic goal is to span the microcrystalline scale and reach grain Fatigue tests were performed in air and in distilled water using the sizes smaller than 100 nm [292]. However, hardly any true bulk cumulative fatigue design approach. Fatigue crack propagation was nanocrystalline metals have been produced due to the difficulty not studied. The main achievement of this work is the verification in producing yields that are sufficiently large to undergo the stan- of a direct relationship between the thickness of the coating layer dardized mechanical tests [293]. and the resistance to corrosion fatigue failure. The authors ob- Because of this, other processes have emerged as viable engi- served that a 3 lm single DLC layer did not enhance the corrosion neering alternatives to circumvent this problem. One example is

Fig. 23. Relative increase in fatigue life as a function of maximum alternating stress for DLC-coated 316L stainless steel in air and in 3 wt.% NaCl solution (from Puchi-Cabrera et al. [277]; reprinted with permission from Elsevier). 956 R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962 the production the so-called ultrafine-grained (UFG) metals with responsible for the increase in the fatigue strength. The authors an average grain size between 100 nm and 1 lm. Today, severe speculate that ECAP can be used to produce Ti–6Al–4V ELI alloy plastic deformation processes, such as equal channel angular for biomedical applications. However, the influence of corrosion pressing (ECAP), are routinely employed to produce bulk submi- on the fatigue properties of the UFG materials was not assessed. cron-grained metallic materials [294]. Wang et al. [301] investigated the corrosion properties of a In addition to the well-known effect of severe plastic deforma- nanocrystalline austenitic stainless steel in a 3.5 wt.% NaCl solution on the mechanical properties of metallic materials, the biolog- tion. A combination of shot peening and heat treatment was used ical response of a UFG alloy may also be positively affected by this to produce surface nanocrystallization. Nanocrystallization was structure modification. An interesting example of this was re- found to decrease the anodic current density and passive current ported by Kim et al. [295]. They showed that ECAP processing en- density of the treated specimens in comparison with the as-re- hanced the biocompatibility of a grade 2 Ti rod in comparison with ceived ones. Furthermore, the surface treatment had a marked a conventional Ti–6Al–4V alloy by providing greater cell prolifera- beneficial effect on the pitting corrosion resistance of the alloy. It tion and adhesion. This behavior was due to the high surface en- greatly extended the passive region of the stainless steel and in- ergy and the presence of nanosized grooves on the severely creased the breakdown potential of the passive film. It is hypothe- plastic deformed material. sized that this effect is related to an increase in the density of Surface nanocrystallization has also arisen as a method to en- diffusion paths available for the migration of alloying elements hance the fatigue properties of metallic materials. The development and rapid formation of a protective passive film in the nanocrystal- of this was motivated by the fact that fatigue failure is very sensitive line surface material. Similar results were reported by Wang and Li to the microstructure and properties of the material’s surface. Thus, [255]. the optimization of the surface microstructure can be effective at Hu et al. [302] studied the corrosion behavior of an orthopedic enhancing the performance of metallic materials under cyclic stres- nanocrystalline NiTi alloy produced by SMAT. They observed that ses [296]. Moreover, other surface-related phenomena, such as the corrosion potential of the nanocrystalline material was shifted wear and corrosion, can also benefit from this approach [297]. Sur- to more positive values in comparison with the untreated alloy. face nanocrystallization can be achieved by surface mechanical Furthermore, the corrosion current density was reduced by one attrition treatment (SMAT) or shot peening [298,299]. The fatigue order of magnitude after the surface treatment. Despite the attrac- properties of biomedical metallic materials have been improved tive indications of these reports from a corrosion standpoint, the by UFG and surface nanocrystallization methods. In this section fatigue behavior of the nanocrystalline materials was not assessed. the mechanisms that govern this enhancement are reviewed. Fur- Roland et al. [303] investigated the fatigue behavior of a surface thermore, the potential of using surface nanocrystallization to mit- nanostructured 316L stainless steel subjected to SMAT. The results igate the corrosion fatigue failure of metallic implants is discussed. obtained from S–N curves showed that the fatigue strength of the Saitova et al. [300] studied the fatigue behavior of a UFG Ti– nanostructured material was significantly higher than that of the 6Al–4V ELI alloy. The ultrafine structure was obtained by combin- untreated specimens in both the low and high cycle fatigue re- ing heat treatment, ECAP and conventional extrusion. Two differ- gimes. Using TEM analysis, the authors could distinguish a nano- ent processing conditions were used to produce the UFG alloys: crystalline surface layer and a plastically deformed sub-surface varying the heat treatment temperature and varying the number layer containing twins and high dislocation densities. The de- of ECAP steps. The materials yielded in the two conditions were formed sub-surface presents high compressive residual stresses, identified as UFG1 and UFG2, respectively. The fatigue behavior while the nanocrystalline layer at the surface accounts for an in- of the UFG materials was compared with that of a conventional crease in mechanical strength by grain size reduction. Under these coarse-grained (CG) alloy through stress-controlled tests in the conditions, fatigue crack initiation and propagation are prevented. high cycle fatigue regime, thus obtaining S–N curves. The results The authors additionally showed that a post-annealing step after are shown in Fig. 24. The fatigue lives of the UFG specimens were SMAT can further enhance the fatigue strength of the material by significantly improved in comparison with the CG ones. The fatigue increasing the surface ductility through a recovery process, with- endurance limit increased by 70 MPa for both UFG conditions. The out leading to grain size growth. The influence of corrosion pro- severe plastic deformation imparted by the ECAP process is cesses on the fatigue behavior of the SMAT-processed material was not assessed. Surface nanocrystallization of magnesium alloys has been successfully achieved by a variety of processes, such as ECAP, SMAT and high-energy shot peening [304–307]. Despite the positive effects that have been reported for the wear resistance and surface microhardness of nanocrystalline magnesium alloys, a detailed investigation of the effect of nanocrystallization on the fatigue properties of these materials is still awaited. Moreover, the corrosion behavior, whether associated with fatigue or not, has not been investigated in the literature. This is especially true for alloys spe- cifically developed for biomedical applications, such as Mg–Ca alloys. From the analysis of the current literature, it can be concluded that nanocrystallization may be a powerful tool for enhancing the fatigue behavior of biomedical metallic materials. Nevertheless, it is imperative to extend the scientific investigations toward a concomitant evaluation of corrosion and fatigue. Only isolated studies of each of these processes have been reported. Moreover, the fatigue crack growth behavior of nanocrystallized biomedical alloys is often ignored in the literature. This approach becomes Fig. 24. S–N curves for Ti–6Al–4V ELI alloy in CG and UFG conditions (from Saitova more important if one considers that severe plastic deformation et al. [300]; reprinted with permission from Elsevier). processes such as ECAP may decrease the fatigue strength of Mg R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962 957 and Ti alloys in the low cycle fatigue regime [296]. The inferior that correlate the mechanisms of corrosion fatigue crack growth of performance of UFG and nanocrystalline metals is related to their biomedical alloys with microstructural features such as grain size, limited ductility under cyclic deformation, which leads to early crystalline phase composition and distribution are very scarce, crack initiation and a greater number of grain boundaries that while experimental investigations that also include different heat are favorably oriented for crack propagation. Even though this treatments and mechanical processing operations are not encoun- problem can be prevented by the combination of ECAP with further tered in the literature at all. mechanical treatments [308,309], it is important to consider early Predicting the fatigue lives of metallic biomaterials in a reliable fatigue crack initiation when designing nanocrystalline biomedical manner will only be possible through the development of unequiv- alloys. In this respect, the assessment of crack growth rate through ocal models based on extensive experimental data. As can be in- da/dN vs. DK curves is of prime interest. This is evident from an ferred from the literature reviewed within this work, this analysis of the results presented in Fig. 25 [240]. The plots in this research field has the potential to be greatly expanded to all the figure were obtained for microcrystalline (mc) and ultrafine-crys- typical metallic materials used for biomedical applications. talline (ufc) grade 2 Ti. It is clear that the DKth values are lower The same scarceness of systematic investigations occurs with for the ufc material. This effect may be related to the grain refine- the methods of preventing and mitigating corrosion fatigue failure. ment and strain hardening of the ECAP-processed material. Conse- Even for traditional techniques such as shot peening and hard thin quently, dislocation mobility is limited and this inevitably results coatings there is an insufficiency of corrosion fatigue studies. This in suppression of the crack tip blunting effect (limited crack tip shortcoming is still more pronounced with the most up-to-date plasticity), leading to faster propagation of fatigue cracks. Micro- methods, which produce UFG or surface nanocrystalline materials. structural changes due to specific alloying additions and heat treat- More than just a fertile research field, corrosion fatigue of biomed- ments may alter this behavior. Systematic studies of such ical metallic materials is also a phenomenon that demands signif- possibilities for different biomedical metallic materials are lacking icant technological advancements to be fully understood and in the literature. prevented.

4. Recommendations for future work 5. Conclusions It is well recognized that corrosion fatigue impairs the performance of biomedical metallic alloys and is still responsible for Mechanical failure of metallic biomaterials is often related to most of the catastrophic failures of these components, despite corrosion fatigue phenomena. Pitting corrosion is often considered the evolution of both material quality and design over the years. to be the active mechanism through which fatigue cracks initiate. Therefore, it would be expected that the development of new The way nucleated cracks propagate depends on the surface and materials and the enhancement of the existing ones would be dri- internal (defects, phases) characteristics of the metallic material. ven by this issue. Indeed, there is a huge amount of information on Microstructural aspects are closely related to the corrosion fatigue the isolated aspects of this problem, i.e. works devoted only to strength of biomedical alloys. This property can be significantly al- corrosion or to fatigue of metallic biomaterials. Many authors have tered through the careful control of alloy composition and by the reported on the combined action of corrosion and fatigue of bio- use of thermo-mechanical processing. The CFD approach is often medical alloys in physiological solutions, facing the challenge of used to characterize the corrosion fatigue behavior of biomedical understanding the synergism between them. However, while alloys. FCP investigations through da/dN vs. DK curves are often ig- cumulative fatigue design has received much attention, fatigue nored in many studies. This common practice yields an incomplete crack propagation has barely been investigated. There is hence a assessment of the corrosion fatigue behavior of metallic biomateri- strong need to investigate the crack growth behavior of metallic als as the presence of pre-existing flaws is very likely and should biomaterials in corrosive physiological environments, composed not be disregarded during the development of a biomedical device. not only of saline species, but also of proteins and enzymes. Studies The lack of investigations devoted to this approach is surprising. Moreover, the contribution of wear processes to the corrosion fatigue failure of orthopedic components such as modular total hip stems has been recognized. This poses a further difficulty to the simulation of actual in vivo conditions during laboratory tests. The complexity of this task is to simultaneously encompass corrosion, fatigue and wear processes of the orthopedic device without ignoring the influence of microstructure and surface finishing on the material response. The most commonly used methods of preventing fatigue failure of metallic materials are based on the introduction of compressive residual stresses or on the deposition of hard thin coatings. Succes- sive examples of these approaches have been outlined throughout the text. Protective coatings may also influence the corrosion resistance of the material, as well as processes that introduce residual stresses. The corrosion fatigue behavior of metallic alloys subjected to different surface modification methods has been studied by several authors. The CFD approach is very frequent, while the FCP approach is not. More recently, UFG or nanocrystalline alloys have emerged as promising materials for biomedical applications. Despite some valuable reports on the suitability of achieving high fatigue strength and corrosion resistance through surface nanocrystallization, the systematic study of this method as a means of Fig. 25. Crack growth rates for grade 2 Ti after ECAP process (from Hanlon et al. improving the corrosion fatigue performance of biomedical metal- [289]; reprinted with permission from Elsevier). lic alloys is only just beginning. 958 R.A. Antunes, M.C.L. de Oliveira / Acta Biomaterialia 8 (2012) 937–962

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