by user

Category: Documents





Acta Biomaterialia 8 (2012) 852–859
Contents lists available at SciVerse ScienceDirect
Acta Biomaterialia
journal homepage: www.elsevier.com/locate/actabiomat
Tribo-electrochemical characterization of metallic biomaterials
for total joint replacement
N. Diomidis a,⇑, S. Mischler a, N.S. More b,1, Manish Roy c
Tribology and Interface Chemistry Group, Swiss Federal Institute of Technology of Lausanne, Lausanne, Switzerland
Department of Metallurgical and Materials Engineering, Visvesvaraya National Institute of Technology, Nagpur, India
Defence Metallurgical Research Laboratory, Hyderabad, India
a r t i c l e
i n f o
Article history:
Received 8 June 2011
Received in revised form 21 September 2011
Accepted 26 September 2011
Available online 29 September 2011
Fretting corrosion
Artificial implants
a b s t r a c t
Knee and hip joint replacement implants involve a sliding contact between the femoral component and
the tibial or acetabular component immersed in body fluids, thus making the metallic parts susceptible to
tribocorrosion. Micro-motions occur at points of fixation leading to debris and ion release by fretting corrosion. b-Titanium alloys are potential biomaterials for joint prostheses due to their biocompatibility and
compatibility with the mechanical properties of bone. The biotribocorrosion behavior of Ti–29Nb–13Ta–
4.6Zr was studied in Hank’s balanced salt solution at open circuit potential and at an applied potential in
the passive region. Reciprocating sliding tribocorrosion tests were carried out against technical grade
ultra high molecular weight polyethylene, while fretting corrosion tests were carried out against alumina.
The wear of the alloy is insignificant when sliding against polyethylene. However, depassivation does
take place, but the tested alloy showed an ability to recover its passive state during sliding. The abrasivity
of the alloy depends on the electrochemical conditions of the contact, while the wear of polyethylene proceeds through third body formation and material transfer. Under fretting corrosion conditions recovery of
the passive state was also achieved. In a fretting contact wear of the alloy proceeds through plastic deformation of the bulk material and wear resistance depends on the electrochemical conditions.
Ó 2011 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction
Human joints operate by low-friction articular cartilage bearing
surfaces, which are conforming and self-regenerating [1–3]. When
natural joints are severely damaged, e.g. due to osteoarthritis, they
are often replaced by artificial implants. In total joint replacement
the implant components are generally made of metal–metal, metal–polymer, ceramic–ceramic or ceramic–polymer couples. Metal
on polyethylene is a very common material coupling in total joint
replacement [4].
Knee and hip replacement joints involve a sliding contact at the
articulation between the femoral component and the tibial or acetabular component during motion of the human body [5–7]. As a
result the metallic components of the artificial joint are susceptible
to sliding tribocorrosion (see Fig. 1). Tribocorrosion is the irreversible transformation of a material due to the simultaneous action of
corrosion and wear taking place in a sliding tribological contact. It
involves numerous synergy effects between mechanical and electrochemical phenomena, usually leading to an acceleration of
material loss [8–10].
⇑ Corresponding author. Tel.: +41 216932952.
E-mail address: [email protected]fl.ch (N. Diomidis).
Present address: Nuclear Power Corporation of India Ltd, Mumbai, India.
Fretting corrosion is a particular form of tribocorrosion involving a small amplitude relative displacement or vibration, usually
between surfaces that are meant to be fixed to each other [11].
In the particular case of orthopedic implants micro-motions are
known to occur at points of fixation [12], while corrosion is caused
by the body fluids, which contain various inorganic and organic
ions and molecules [13] (see Fig. 1). Fretting corrosion has been
identified at the stem/neck and neck/head contacts of modular implants, at the stem/bone and stem cement interfaces of cemented
and uncemented implants, and at the screw/plate junction of fixation plates [14]. In contrast to sliding, during fretting a considerable part of the displacement may be accommodated in the
contact by elastic deformation and thus the elastic properties of
biomaterials can affect the implants behavior and functionality.
Furthermore, due to the closed geometry of the contact electrolyte
replenishment is difficult and the electrochemical conditions
might differ from those of the bulk. Furthermore, debris is easily
trapped and thus the behavior of the third body is critical [15].
The commonly used metallic biomaterials for orthopedic applications (stainless steel, titanium and CoCrMo alloys) owe their high
corrosion resistance to the spontaneous formation of a passive
surface oxide layer which forms the interface between the alloy
and the environment [16,17]. The properties of surface films dictate the results of chemical and mechanical interactions at the
1742-7061/$ - see front matter Ó 2011 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
N. Diomidis et al. / Acta Biomaterialia 8 (2012) 852–859
Fig. 1. A schematical representation of a total hip joint replacement prosthesis. The
types of motion and surface degradation mechanisms of the implant metallic
components are shown.
interface. Both sliding tribocorrosion and fretting corrosion of passive metals lead to local damage or removal of the passive film, as
well as detachment of metal particles, leading to mechanical wear.
As the counterbody moves on the depassivated surface area
re-oxidizes, a process involving a charge transfer reaction at the
interface which yields dissolved metal ions and a solid oxide. As
a result in both tribocorrosion situations passive film degradation
induces wear-accelerated corrosion. The repeated removal of oxide
films produces particles and ions, which can result in adverse
biological reactions and can lead to mechanical failure of the device [18–21]. Therefore, an approach combining both electrochemistry and tribology is a necessary means to study these complex
phenomena and to assess the biocompatibility of candidate metallic materials.
In recent years different electrochemical techniques have been
combined with tribology [22] and methodologies have been developed [23] that allow the study of tribo-electrochemical systems.
Two types of electrochemical tests are commonly combined with
tribology: chronopotentiometry and potentiostatic polarization.
During chronopotentiometric measurements the evolution of the
open circuit potential (Eoc) of the sample is monitored before,
during and after a mechanical perturbation. The open circuit
potential is the difference in electrical potential spontaneously
established between the metal and the solution. Upon disruption
of the passive surface film the underlying metal is uncovered and
the Eoc drops to more cathodic values, indicating the increased oxidation tendency of the exposed metal. In such a case the measured
Eoc is a mixed potential resulting from galvanic coupling between
passive and depassivated areas in the sample surface [24]. On the
other hand, during potentiostatic polarization tests one can control
the potential and thus study the response of the material under different oxidizing conditions, such as in the case of local tissue
inflammation when oxidizing agents will be present in the environment. The potential of the sample is externally imposed and
the evolution of the current is monitored during sliding or fretting.
In this case disruption of the passive surface films leads to an increase in the measured anodic current originating from the accelerated oxidation reactions taking place at the uncovered metal
surface [25].
The scope of this work is to evaluate the combination of electrochemistry and tribology in the study of the wear behavior of biomedical alloys. For this, laboratory set-ups are used to test a
newly developed b-titanium alloy in different tribological contacts
in physiological solution. The selected counterparts were
UHMWPE and alumina in order to simulate the ball/acetabular
cup sliding contact and the femoral stem/ball fretting contact,
2. Materials and methods
2.1. Characteristics of the alloy
In this work a titanium alloy with b microstructure (see Fig. 2),
Ti–29Nb–13Ta–4.6Zr, was studied [26,27]. b-Alloys exhibit a modulus of elasticity considerably lower than commonly used Ti–6Al–
4V and Co–28Cr–6Mo alloys and thus decrease the risk of stress
shielding of the femur [28]. The added benefit of using Nb, Zr
and Ta as alloying elements is that they tend to form dense surface
oxides, increasing the stability of the passive layer [29].
The alloy was prepared by melting in a vacuum arc electric furnace (non-consumable) and casting in the form of pancakes, followed by forging. A vacuum arc electric furnace in which the
charge can be melted in a water cooled copper crucible under vacuum (103 mbar) was employed. Prior to melting the furnace
chamber was purged with argon gas once and then again filled
with the gas up to 532 mbar. The charge was melted by arcing
emitted from a tungsten bit brazed to a copper stringer rod suspended above the charge. A d.c. potential of up to 30–32 V and current of 1000 A was applied between the tungsten cathode and
charge material, which formed the anode. A stirring coil around
Fig. 2. (a) Optical micrograph and (b) X-ray diffraction spectrum of the Ti–29Nb–13Ta–4.6Zr alloy.
N. Diomidis et al. / Acta Biomaterialia 8 (2012) 852–859
the copper crucible stirred the melt pool and homogenized the
melt composition. Each pancake was melted four times to produce
a homogeneous composition. The pancake was then forged at a
temperature of 900 °C in a forging press. The composition of the alloy measured by inductively coupled plasma optical emission
spectrophotometry is 27.2 wt.% Nb, 11.5 wt.% Ta, 4.8 wt.% Zr,
0.05 wt.% S, 0.05 wt.% C, and 230 ppm H2. The Ti–29Nb–13Ta–
4.6Zr alloy has a hardness of 249 MPa and an elastic modulus of
55 GPa.
Cylindrical specimens were cut with dimensions of 20 mm
diameter and 5 mm thickness. Sample preparation was done by
progressively grinding with emery papers of 240, 400 and 500 grit,
followed by fine polishing with diamond paste (6 lm) and final
polishing with alumina powder (3 lm), leading to an Ra of
0.05 lm. The specimens were then washed in deionized water
and ultrasonically cleaned in acetone and ethanol for 10 min each.
2.2. Experimental conditions
All tests were done in Hank’s balanced salt solution (HBSS), consisting of NaCl (8.00 g l1), KCl (0.40 g l1), CaCl2 (0.18 g l1), NaHCO3 (0.35 g l1), NaHPO42H2O (0.48 g l1), and MgCl26H2O
(0.10 g l1) at 37 ± 1 °C. A Wenking LB 95 L Auto Range laboratory
Potentiostat served to control the potential of the titanium specimen disk (working electrode). A platinum wire and a mercury sulfate electrode (MSE) were connected to the potentiostat as counter
and reference electrodes, respectively. Experiments were performed at open circuit potential (Eoc) and at an applied potential
of +2 V vs. MSE corresponding to the passive region (Epass), as
shown in Fig. 3. For tests at open circuit potential the samples were
immersed in the electrolyte 1800 s before the initiation of sliding.
For tests at a passive applied potential the samples were immersed
in the solution at rest potential for 300 s and then the anodic potential was applied for 1800 s prior to the initiation of sliding.
Reciprocating sliding tribocorrosion tests (Fig. 4) were carried
out with a sliding tribo-electrochemical apparatus, details of which
are given in a separate publication [30]. Pins with a hemispherical
end of 12 mm diameter and 38 mm length were made up of technical grade UHMWPE and were used as counterbodies. The applied
load was 6.5 N, at a frequency of 1 Hz, with a stroke length of 5 mm
for 3600 s. This resulted in an initial maximum Hertzian contact
pressure of 23 MPa. The wear of the hemispherical tip of the polymer pin was determined by measuring the flat end formed during
rubbing on optical micrographs and geometrically calculating the
corresponding volume of the worn spherical cup.
Fretting corrosion tests (Fig. 4) were carried out with a fretting
tribo-electrochemical apparatus described in detail elsewhere [31].
Alumina balls of 10 mm diameter (SWIP AGBrügg, G10 AFBMA finish) were used as counterbodies. Fretting corrosion experiments
were carried out at an applied normal load of 10 N, with a displacement of 100 lm and a frequency of 1 Hz was applied for 3600 s.
This resulted in an initial maximum Hertzian contact pressure of
400 MPa. The vertical position of the counterbody was monitored
using a Keyence LC2420 laser distance meter at a resolution of
0.01 lm. The mean frictional coefficient (l) was calculated by
dividing the tangential force by the normal force when the ball
was in the middle of the stroke. Laser profilometer (UBM Telefokus
UBC14) surface scans were performed at a resolution of
300 points mm1 and the wear volumes were calculated according
to Fouvry et al. [32]. Scanning electron microscopy (Philips XL30FEG SEM) and optical microscopy were used to characterize the
morphology of the wear tracks.
3. Results
3.1. Sliding tribocorrosion
The evolution of the coefficient of friction and the potential with
time during sliding tribocorrosion experiments on Ti–29Nb–13Ta–
4.6Zr at open circuit potential are shown in Fig. 5. The coefficient of
friction goes through a run-in period of about 700 cycles and then a
steady-state value of around 0.3 is achieved. Before the initiation of
sliding the measured open circuit potential reflects the presence of
a passive film on the alloy surface in contact with the electrolyte
Fig. 3. Electrochemical potentio-dynamic polarization of Ti–29Nb–13Ta–4.6Zr in
Fig. 4. Schematic diagram of the tribological test systems.
Fig. 5. Evolution of the coefficient of friction and the open circuit potential during
sliding tribocorrosion experiments at Eoc.
N. Diomidis et al. / Acta Biomaterialia 8 (2012) 852–859
Table 1
The wear volume of UHMWPE pins after sliding tribocorrosion experiments on Ti–
29Nb–13Ta–4.6Zr at open circuit potential and at a passive applied potential.
Electrochemical condition
Wear volume (104 mm3)
4.08 ± 1.12
1.62 ± 0.16
transfer film is detected at open circuit potential (Fig. 7), indicating
that polyethylene transfer is not responsible for the measured electrochemical response. The wear of the UHMWPE pin after tribocorrosion tests at different potentials is shown in Table 1.
3.2. Fretting corrosion
Fig. 6. Evolution of the coefficient of friction and the anodic current, during
tribocorrosion experiments at a passive applied potential. The dotted line indicates
the background current in the absence of sliding, calculated according to Mischler
et al. [33].
[24]. Immediately upon the initiation of sliding the open circuit potential decreases, indicating a depassivation of the surface induced
by mechanical removal of the passive film. For a certain period of
time during sliding the open circuit potential remains around that
low value, indicating that a depassivated state prevails on a part of
the surface. During this period the potential exhibits slight variations of about 75 mV resulting from local depassivation–repassivation phenomena. After about 700 s of sliding the open circuit
potential suddenly increases until it reaches values similar to that
measured before the initiation of sliding. This is due to regrowth of
the passive film in the wear track, indicating that at open circuit
the alloy has the ability to recover its passive state while mechanical perturbation is taking place.
The evolution of the coefficient of friction and the anodic current (i) with time during sliding tribocorrosion experiments on
Ti–29Nb–13Ta–4.6Zr at a passive applied potential are shown in
Fig. 6. At Epass a steady-state coefficient of friction value of around
0.3 is achieved after a run-in period. An increased anodic current is
measured at the onset of sliding as a result of depassivation. During
the course of the experiment a considerably increased current is
measured during running in. Then it gradually decreases until
the measured current is of the same order as that expected without
sliding. This indicates that the material has the ability to regain its
passive state during sliding at Epass.
After sliding tribocorrosion experiments both the alloy and the
UHMWPE counterbody were examined for wear. No measurable
wear was found on the metallic alloy samples by non-contact profilometry. A surface microstructure characteristic of polyethylene
transfer covering a part of the wear track is found at Epass, but no
Acquisition of transient values during fretting allows the plotting of fretting logs, i.e. three-dimensional graphical representations of the time evolution of the frictional force–displacement
loops. Such diagrams are shown in Fig. 8 for fretting corrosion
experiments carried out at open circuit and a passive potential.
The fretting log diagrams have an open trapezoidal shape, indicating that fretting follows a gross slip regime. Both elastic deformation and slip at the interface contribute to accommodation of the
imposed displacement. Elastic accommodation of 26 and 32 lm
of the imposed 100 lm occurs at Eoc and Epass, respectively.
The evolution of the coefficient of friction and potential with
time during fretting corrosion experiments on Ti–29Nb–13Ta–
4.6Zr at open circuit potential are shown in Fig. 9. A steady-state
coefficient of friction value of 0.65 is measured during fretting at
Eoc. Regarding the evolution of the open circuit potential, a similar
response to that measured under sliding tribocorrosion conditions
is revealed. The open circuit potential decreases upon the initiation
of fretting. However, the potential drop (60 mV) is smaller than
the potential drop measured upon the initiation of sliding
(200 mV), since the depassivated area under fretting is smaller
than under sliding due to the considerably smaller imposed displacement (100 lm vs. 5 mm). After only a few hundred cycles it
rises again to reach values similar to those measured before the onset of fretting. Thus passive state recovery is achieved during fretting corrosion experiments at Eoc.
The evolution of the coefficient of friction, the anodic current
and the vertical position of the counterbody with time during fretting corrosion experiments on Ti–29Nb–13Ta–4.6Zr at a passive
applied potential are shown in Fig. 10. After a run-in period a steady-state coefficient of friction value of around 0.6 is measured during fretting at Epass. For fretting corrosion tests at Epass the anodic
current rises immediately on the initiation of fretting, indicating
depassivation in the wear track. After about 2500 s of fretting the
excess current decreases, indicating passive state recovery in the
wear track. Thus passive state recovery is achieved during fretting
Fig. 7. Scanning electron micrographs of the wear tracks on Ti–29Nb–13Ta–4.6Zr after tribocorrosion testing at (a) Eoc and (b) Epass.
N. Diomidis et al. / Acta Biomaterialia 8 (2012) 852–859
Fig. 8. Fretting log diagrams of Ti–29Nb–13Ta–4.6Zr at (a) open circuit potential and (b) a passive applied potential.
value when indentation stops and passivity is recovered, indicating
a close correlation between the mechanical, electrochemical and
wear phenomena in the contact.
After the fretting corrosion experiments both the alloy and the
alumina counterbody were examined for wear. No measureable
wear of the alumina was found, with only a few debris particles
from the metal being found adhering to the ball. In contrast, wear
does take place on the titanium alloy samples at both open circuit
potential and at a passive potential. At both potentials a similar
microstructure of the wear track largely characterized by plastic
deformation and mixing of the third body is obtained (Fig. 11). A
large number of debris particles were found surrounding the wear
track. The wear of the alloy after fretting corrosion tests at different
potentials is shown in Table 2.
4. Discussion
Fig. 9. Evolution of the coefficient of friction and the potential during fretting
corrosion experiments at open circuit potential.
4.1. Influence of electrochemical conditions on overall wear
at Epass. From Fig. 10 it can be seen that the excess current during
fretting is measured only while the counterbody moves downwards, indenting the test sample. The excess current decreases to
almost zero at the end of indentation of the sample. As a result,
recovery of the passive state on the sample surface during fretting
is accompanied by zero wear. Interestingly, the run-in period of the
coefficient of friction during which large variations are measured
coincides with the duration of indentation of the sample by the
counterbody. The coefficient of friction reaches a steady-state
When sliding is imposed on a Ti–29Nb–13Ta–4.6Zr/UHMWPE
contact wear of the polymer occurs but no measurable wear of
the alloy is found. The wear of the polymer differs depending on
the imposed potential, indicating that the abrasivity of the alloy
is dependent on the oxidative conditions prevailing in the contact
zone. The wear volume of the polymer pin at Eoc is considerably
larger that the wear volume measured after sliding tests at Epass.
Additionally, the scanning electron microscopic examination indicated the presence of a polyethylene transfer film at Epass but not at
Eoc. As a result a three-body contact exists at Epass, in contrast to
Fig. 10. Evolution of the coefficient of friction, the anodic current and the vertical position of the counterbody during fretting corrosion experiments in Ti–29Nb–13Ta–4.6Zr
at a passive applied potential. The dotted line indicates the background current in the absence of fretting, calculated according to Mischler et al. [33].
N. Diomidis et al. / Acta Biomaterialia 8 (2012) 852–859
Fig. 11. Representative scanning electron micrograph of the wear track after
fretting corrosion of Ti–29Nb–13Ta–4.6Zr at Epass.
Table 2
The wear volume of the Ti–29Nb–13Ta–4.6Zr samples after fretting corrosion
experiments at open circuit potential and at a passive applied potential.
Electrochemical conditions
Wear volume (104 mm3)
9.87 ± 1.02
4.35 ± 0.02
depth of the material lost due to chemical wear is 43 nm. In reality
this depth is expected to be even smaller, since the contact area and
thus the wear track width increase in size with wear of the
polyethylene pin. This confirms that the material lost due to
wear-accelerated corrosion is indeed very small and comparable
with the initial surface roughness, as was seen from microstructural
examination of the surface.
Applying a similar calculation to the fretting corrosion data of
Fig. 10 a 9 106 mm3 volume of metal is lost due to wearaccelerated corrosion. This represents only about 2% of the total
worn volume, or an increase in the depth of the wear track of about
0.1 nm. This indicates that the mechanical character of wear predominates during fretting, in accordance with the smaller wear
track area and higher contact pressure compared with sliding
tribocorrosion tests.
The above mentioned calculation highlights the different prevailing material removal mechanisms depending on the tribological contact. Under sliding conditions an electrochemical
oxidation degradation mechanism predominates, accelerated by
mechanical removal of the protective corrosion products. Under
fretting mechanical removal of bulk material is the dominant degradation mechanism.
4.3. Recovery of the passive state
Eoc, at which an alloy–polymer two-body contact predominates.
The above mentioned results reveal that the presence of the
UHMWPE transfer film, which protects the polyethylene pin from
further wear, is influenced by the electrochemical conditions in
the contact.
In the case of the fretting corrosion experiments a different situation prevails. This is not surprising since alumina is much harder
than the metal while UHMWPE is softer. Under fretting the alumina counterbody shows no indication of wear, while measurable
wear of the Ti alloy is found. In this case also the electrochemical
conditions influence the total wear volume. A smaller amount of
wear is measured after tests done at a passive potential than at
Eoc. This could be explained by the different structure and chemical
composition, as well as the greater thickness of the passive layer at
higher potentials [29,34,35]. However, after a certain period of
fretting wear of the alloy stops.
4.2. Wear-accelerated corrosion
Even though no wear of the metallic alloy can be measured, the
electrochemical parameters (potential and current) measured during sliding reveal that depassivation of the surface does take place.
Depassivation will lead to material loss since the uncovered metallic surface will spontaneously re-oxidize, releasing metal ions,
indicating synergy between mechanical and electrochemical phenomena. In the case of tests at an applied potential the excess
current due to sliding can be measured (see Fig. 6) and the amount
of oxidized material can be calculated using Faraday’s law:
where I is the excess current due to rubbing, t is the duration of
rubbing, M is the atomic mass of the metal, n is the metal oxidation
valence, F is the Faraday constant, and d is the density. For tests
done at open circuit potential such a calculation cannot be made
because no external current circulates. If the measured excess
current is assumed to result from the oxidation of Ti to TiO2 a total
of 1.61 104 mm3 of metallic material is lost due to wearaccelerated corrosion. Assuming that the size of the wear track
can be approximated by multiplying the initial Hertzian contact
diameter (740 lm) with the stroke length (5 mm), then the average
When sliding tribocorrosion experiments are carried out at
either open circuit or applied potential the electrochemical response of the alloy during sliding reaches values expected without
sliding. This indicates that the alloy tends to regain its passive state
despite mechanical perturbation. The ability of an alloy to regain
its passive state while sliding is a critical property in orthopedic
prostheses, since the release of metallic ions due to wear-accelerated corrosion will be limited during the lifetime of the implant.
For a material to be able to recover its passive state during sliding
after depassivation has taken place it is necessary that depassivation ceases, since the repassivation rate of the metal does not
change considerably during the course of the experiment. In order
for this to happen the stress acting on the passive surface needs to
be lower than the critical value required for passive film removal.
Thus the nature of the passive film and particularly adhesion to the
substrate and the ability to resist delamination are critical properties. In the present experiments the friction coefficient is stable and
does not change considerably with potential. Thus the critical
parameter for passive film removal is the contact pressure. Since
the metallic alloy shows very little wear the evolution of the contact pressure from the initially applied 23 MPa during the course of
the experiment depends on wear of the polyethylene. For sliding
tests carried out at Eoc, at which a thin passive layer is expected,
depassivation takes place immediately upon initiation of sliding.
On the other hand, for tests at Epass, at which a thicker passive film
is expected [29], apart from an initial current peak, the anodic current does not increase significantly until a few hundred cycles have
occurred. This indicates that the thickness of the passive layer is
also a critical parameter for depassivation.
In the case of fretting corrosion an initial maximum contact
pressure of 400 MPa is applied. During the course of the test only
the metallic alloy sample wears, while the alumina counterbody
is unaffected. As a result, the evolution of the pressure in the
contact depends on wear of the alloy. The link between the wear
volume and the recovery of passivity has been demonstrated for
a b-Ti–13Nb–13Zr alloy [27]. According to Fig. 10 depassivation
of the alloy stops at the moment when wear also stops. Since wear
of the alloy stops the size of the wear track measured at the end of
the test can be assumed to be the same as at the moment of
passivity recovery. Measuring the size of the wear track allows
approximate calculation of the pressure at the moment of passivity
N. Diomidis et al. / Acta Biomaterialia 8 (2012) 852–859
of conditions where corrosion is significant or not. Such phenomena have a high clinical relevance, and thus the tribo-electrochemical approach could be beneficially applied to
conditions more representative of the in vivo situation (e.g.
hip and knee simulators).
Tribo-electrochemical techniques have shown that Ti–29Nb–
13Ta–4.6Zr recovers a passive surface state under both sliding
and fretting contacts in a physiological solution. This is a critical
property for biomedical applications, which if taken into
account in alloy development and implant design could lead
to decreased material loss and increased biocompatibility.
Fig. 12. Schematic representation of the loading conditions and the depassivation
mechanism during sliding and fretting tests.
recovery. The critical pressure for depassivation in the different
fretting corrosion tests is in the range 100–200 Mpa, as shown in
Fig. 12. Such a pressure is considerably larger than the critical
pressure for passive film removal found under sliding, which is
lower than 23 MPa.
This difference in critical pressure for depassivation can be
attributed to different depassivation mechanisms. Under sliding
no wear or plastic deformation of the alloy takes place according
to Fig. 7. Depassivation during the early stages of sliding proceeds
by passive film delamination. On the other hand, under fretting the
application of a contact pressure higher than the hardness of the
alloy results in plastic deformation of the metallic material and
in the production of third body particles. This plastic deformation
of the material in the wear track is responsible for depassivation.
As fretting proceeds mechanical mixing and compaction of the surface oxides takes place. Furthermore, plastic deformation induces
the formation of a nanocrystalline tribologically transformed structure in the immediate subsurface which is considerably harder
than the base alloy [36]. The presence of such a structure could
be confirmed by nanoindentation measurements.
The phenomenon of passivity recovery during sliding and fretting discussed above appears to be a characteristic of the alloy.
Other b-titanium alloys have exhibited the capability to recover a
passive state under either sliding or fretting conditions [26,27].
However, only Ti–29Nb–13Ta–4.6Zr consistently recovered passivity under both sliding and fretting. The typically used Ti6Al4V does
not recover passivity when tested in the same fretting tribometer
under similar conditions [11,14].
5. Conclusions
A tribo-electrochemical approach has been applied to the study
of the biotribocorrosion behavior of a newly developed b-titanium
alloy for biomedical applications in simulated body fluid in fretting
and sliding contacts. The combination of electrochemistry and tribology has been demonstrated to constitute a powerful tool for the
characterization of biomaterials for joint replacement.
Electrochemical conditions are crucial in the wear of materials
for biomedical applications.
Under sliding the electrochemical conditions can influence the
behavior of the third body and the build-up of polyethylene
transfer films.
Under fretting electrochemical parameters can be used as an
in situ measure of wear.
Tribo-electrochemical tests can quantify the amount of depassivation and the abrasivity of materials.
Combining tribology with electrochemistry allows monitoring
of passive film removal and regrowth, and the identification
Appendix A:. Figures with essential color discrimination
Certain figures in this article, particularly Figs. 1, 3–6, 9, 10 and
12, are difficult to interpret in black and white. The full color
images can be found in the on-line version, at doi:10.1016/
[1] Podsiadlo P, Kuster M, Stachowiak GW. Numerical analysis of wear particles
from non-arthritic and osteoarthritic human knee joints. Wear
[2] Dowson D. A comparative study of the performance of metallic and ceramic
femoral head components in total replacement hip joints. Wear
[3] Podsiadlo P, Stachowiak GW. 3-D imaging of surface topography of wear
particles found in synovial joints. Wear 1999;230:184–93.
[4] Wilches LV, Uribe JA, Toro A. Wear of materials used for artificial joints in total
hip replacements. Wear 2008;265:143–9.
[5] Long M, Rack HJ. Titanium alloys in total joint replacement – a materials
science perspective. Biomaterials 1998;19:1621–39.
[6] Jasty M. Clinical reviews: particulate debris and failure of total hip
replacements. J Appl Biomater 1993;4:273–6.
[7] Niinomi M. Recent metallic materials for biomedical applications. Metall
Mater Trans A 2002;33:477–86.
[8] Diomidis N, Celis JP, Ponthiaux P, Wenger F. Tribocorrosion of stainless steel in
sulfuric acid: identification of corrosion-wear components and effect of
contact area. Wear 2010;269:93–103.
[9] Landolt D, Mischler S, Stemp M, Barril S. Third body effects and material fluxes
in tribocorrosion systems involving a sliding contact. Wear 2004;256:517–24.
[10] Iwabuchi A, Lee JW, Uchidate M. Synergistic effect of fretting wear and sliding
wear of Co-alloy and Ti-alloy in Hank’s solution. Wear 2007;263:492–500.
[11] Barril S, Mischler S, Landolt D. Electrochemical effects on the fretting corrosion
behaviour of Ti6Al4V in 0.9% sodium chloride solution. Wear
[12] Windler M, Klabunde R. Titanium for hip and knee prostheses. In: Brunette
DM, Tengvall P, Textor M, Thomsen P, editors. Titanium in
Medicine. Berlin: Springer-Verlag; 2001. p. 703–46.
[13] Hallab NJ, Jacobs JJ. Orthopedic implant fretting corrosion. Corros Rev
[14] Hiromoto S, Mischler S. The influence of proteins on the fretting-corrosion
behaviour of a Ti6Al4V alloy. Wear 2006;261:1002–11.
[15] Diomidis N, Mischler S. Third body effects on friction and wear during the
fretting of steel contacts. Tribol Int: in press. doi:10.1016/j.triboint.2011.
[16] Virtanen S, Milosev I, Gomez-Barrena E, Trebse R, Salo J, Konttinen YT. Special
modes of corrosion under physiological and simulated physiological
conditions. Acta Biomater 2008;4:468–76.
[17] Milosev I. Metallic materials for biomedical applications: laboratory and
clinical studies. Pure Appl Chem 2011;83:309–24.
[18] Jones FH. Teeth and bones: applications of surface science to dental materials
and related biomaterials. Surf Sci Rep 2001;42:75–205.
[19] Marino CEB, Mascaro LH. EIS characterization of a Ti-dental implant in
artificial saliva media: dissolution process of the oxide barrier. J Electroanal
Chem 2004;568:115–20.
[20] Hodgson AWE, Mueller Y, Forster D, Virtanen S. Electrochemical
characterization of passive films on Ti alloys under simulated biological
conditions. Electrochim Acta 2002;47:1913–23.
[21] Merritt K, Brown SA. Biological effects of corrosion products from metals. In:
Fraker AC, Griffin CD, editors. Corrosion and degradation of implant materials,
STP 859. West Conshohocken, PA: ASTM International; 1983. p. 195–207.
[22] Mischler S. Triboelectrochemical techniques and interpretation methods in
tribocorrosion: a comparative evaluation. Tribol Int 2008;41:573–83.
[23] Diomidis N, Celis JP, Ponthiaux P, Wenger F. A methodology for the assessment
of the tribocorrosion of passivating metallic materials. Lubr Sci
N. Diomidis et al. / Acta Biomaterialia 8 (2012) 852–859
[24] Diomidis N, Göckan N, Ponthiaux P, Wenger F, Celis JP. Assessment of the
surface state behavior of Al71Cu10Fe9Cr10 and Al3Mg2 complex metallic alloys
in sliding contacts. Intermetallics 2009;17:930–7.
[25] Stemp M, Mischler S, Landolt D. The effect of mechanical and electrochemical
parameters on the tribocorrosion rate of stainless steel in sulphuric acid. Wear
[26] More NS, Diomidis N, Paul SN, Roy M, Mischler S. Tribocorrosion behavior of b
titanium alloys in physiological solutions containing synovial components.
Mater Sci Eng C 2011;31:400–8.
[27] Diomidis N, Mischler S, More NS, Roy M, Paul SN. Fretting-corrosion
behavior of b titanium alloys in simulated synovial fluid. Wear 2011;271:
[28] Pilliar RM. Metallic biomaterials. In: Narayan R, editor. Biomedical
materials. Berlin: Springer Verlag; 2009. p. 41–81.
[29] Diomidis N. The surface degradation of metal joints. In: Affatato S, editor. Wear
of Orthopaedic Implants and Artificial Joints. Cambridge: Woodhead
Publishing, in press.
[30] Stojadinović J, Bouvet D, Declercq M, Mischler S. Effect of electrode potential
on the tribocorrosion of tungsten. Tribol Int 2009;42:575–83.
[31] Barril S, Debaud N, Mischler S, Landolt D. A tribo-electrochemical apparatus for
in vitro investigation of fretting-corrosion of metallic implant materials. Wear
[32] Fouvry S, Kapsa Ph, Zahouani H, Vincent L. Wear analysis in fretting of hard
coatings through a dissipated energy concept. Wear 1997;203(/204):393–493.
[33] Mischler S, Barril S, Landolt D. Fretting corrosion behavior of Ti-6Al-4V/PMMA
contact in simulated body fluid. Tribology 2009;3:16–23.
[34] Milosev I, Metikos-Hukovic M, Strehblow HH. Passive film on orthopaedic
TiAlV alloy formed in physiological solution investigated by X-ray
photoelectron spectroscopy. Biomaterials 2000;21:2103–13.
[35] Karthega M, Raman V, Rajendran N. Influence of potential on the
electrochemical behaviour of b titanium alloys in Hank’s solution. Acta
Biomaterialia 2007;3:1019–23.
[36] Sauger E, Fouvry S, Ponsonnet L, Kapsa Ph, Martin JM, Vincent L. Tribologically
transformed structure in fretting. Wear 2000;245:39–52.
Fly UP