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178 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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Mechanical integrity
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of magnesium alloys
5
for biomedical applications
Lokesh Choudhary, R.K. Singh Raman
Monash University, VIC, Australia
5.1 Magnesium and its alloys as implant materials
Human longevity considerably improved in the last century with significant advances
in medical science. The most notable improvements include development of implant
devices, since clinical cases requiring medical implants are increasing with desire
for an active lifestyle and ageing population. Implants are the artificial devices used
in the human body for repairing fractured bones and restoration of any impaired
physical functions (Niinomi, 2002; Ratner, Hoffman, Schoen, & Lemons, 1996).
Common use of implant devices can be varied based on the requirements, such as:
(1) to aid healing process by using plates, rods, and screws for stabilization of fractured
bones, (2) to improve the function of an organ or other parts of the body, and (3) to
replace a damaged, diseased, or worn part of the anatomy. Since implant devices
are used in the physiological environment, they must satisfy certain requirements,
which include excellent biocompatibility, sufficient mechanical properties to hold or
support the injured body part/tissues, and desired resistance to degradation due to
corrosion and fatigue in aggressive human body fluid (Grimm, 2007, pp.
1165e1194; Nielsen, 1987).
In the last two decades, several research groups have investigated the idea that
biodegradable metallic materials might be used for construction of implant devices.
These temporary biomaterials will support tissue regeneration and healing process
and then completely degrade away harmlessly. Traditional implants of stainless steels,
cobaltechromium alloys, and titanium alloys have been widely used because of their
strength, ductility, and resistance to corrosion (Staiger, Pietak, Huadmai, & Dias,
2006). However, when these traditional alloys are used as temporary implant devices,
such as plates, screws, and wires, a second surgery is required to remove the implant
after tissues have healed. Besides the cost of these surgical procedures and inconvenience to patients, the traditional alloys also cause local inflammation due to potential
release of cytotoxic ions (Granchi et al., 1999; Puleo & Huh, 1995; Wang et al., 2002).
Therefore, it would be very advantageous if an implant material could be identified that
degrades away in the physiological environment after completion of the healing
process (Staiger et al., 2006; Witte et al., 2009). This approach will eliminate the
need for the second surgical procedure. In this context, magnesium (Mg)-based
Surface Modification of Magnesium and its Alloys for Biomedical Applications. http://dx.doi.org/10.1016/B978-1-78242-077-4.00005-X
Copyright © 2015 Elsevier Ltd. All rights reserved.
180 Surface Modification of Magnesium and its Alloys for Biomedical Applications
metallic alloys are potential candidates for use as biodegradable temporary implant
devices because they can naturally degrade in the human body environment due to
their high electrochemical activity (standard reduction potential of Mg is 2.4 V vs
standard hydrogen electrode (SHE) (Makar & Kruger, 1993)). Mg, which is nontoxic
to the human body, is also vital for human metabolic processes, including for the
stabilization of DNA and RNA (Hartwig, 2001; Saris, Mervaala, Karppanen, Khawaja,
& Lewenstam, 2000; Wolf & Cittadini, 2003). Any excess Mg is harmlessly excreted
with urine. Unlike traditional implant materials, the mechanical properties of Mg and
its alloys are also very suitable for implant applications: they possess low density
(r) ¼ 1.74e2.0 g cm
3
and elastic modulus (E) ¼ 41e45 GPa, both of which are
similar to the properties of human bones (Staiger et al., 2006). Hence the problem
of stress shielding, which is caused by a mismatch in the elastic modulus between
natural bone and implant, can be mitigated if Mg alloys are deployed as implant
materials (Staiger et al., 2006; Witte, 2011; Xin, Hu, & Chu, 2011). Despite these highly advantageous properties of Mg, its alloys have rarely been used as human body
implants. The major drawback in the use of Mg alloys as implants is that they tend
to corrode very quickly in the corrosive human body fluid (that has a
pH ¼ 7.4e7.6) (Bobby Kannan & Singh Raman, 2008; Heublein et al., 2003;
Kirkland, Lespagnol, Birbilis, & Staiger, 2010; Kirkland, Staiger, Nisbet, Davies, &
Birbilis, 2011; Song, 2007; Staiger et al., 2006; Xin et al., 2011), thereby losing their
mechanical integrity before the expected service life. Although these limitations have
ruled out any effective use of Mg alloys as permanent implants, it is interesting that the
Mg alloys could still be used as biodegradable temporary implant devices such as
plates, wires, stents, pins, and screws.
5.2 Corrosion and assisted cracking of magnesium alloys
Mg and its alloys have been reported to suffer considerably rapid corrosion in the body
environment that contains aggressive ions such as chlorides, carbonates, phosphates,
and sulphates, in addition to the presence of proteins and glucose (Xin, Huo, Tao,
Tang, & Chu, 2008). When used as biomaterials, Mg alloys will be susceptible to a
few different forms of corrosion.
5.2.1 Galvanic corrosion
When two metals with different electrochemical potentials are in physical contact
(electrically) and are exposed to an ion-conducting fluid, such as serum or physiological solution, galvanic corrosion will occur (Jones, 1996). Biodegradability is the
major advantage in using Mg alloys as temporary implants. However, the problem
of severe galvanic corrosion is one of the major obstacles in the successful use of
Mg alloys as implants. Mg makes an active anode when in contact with most metals
because of its high electrochemical activity, and it corrodes preferentially (Makar &
Kruger, 1993; Song & Atrens, 1999).
Mechanical integrity of magnesium alloys for biomedical applications 181
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Mg alloys are also not uniform in terms of their composition, microstructure, and
even crystalline orientation. These differences can result in various electrochemical
activities within a Mg alloy, thereby generating microgalvanic couples. The following
two factors generally influence the formation of micro-galvanic cells in Mg alloys.
5.2.1.1 Impurities
Mg alloys contain iron (Fe), nickel (Ni), cobalt (Co), and copper (Cu) as impurities
(Baker & Avedesian, 1999; Makar & Kruger, 1993; Song & Atrens, 1999) that are
usually introduced during casting processes. These metallic impurities are highly
cathodic to Mg and cause severe galvanic corrosion (Makar & Kruger, 1993; Song
& Atrens, 1999), even when present in small amounts (as shown in Figure 5.1 (Makar
& Kruger, 1993)).
Decreasing contents of these impurities below threshold levels can dramatically
improve the alloy’s corrosion resistance (Makar & Kruger, 1993). When these impurity
contents are below critical concentrations (e.g., their respective solubility in the matrix
phase), they are present in the form of solutes in Mg solid solutions, and therefore, no
microgalvanic cells form between the impurities and Mg matrix. The tolerance limits
for Fe, Cu, and Ni are 170, 1000, and 5 ppm, respectively (Makar & Kruger, 1993).
5.2.1.2 Secondary phases
Most of secondary phases that form in Mg alloys are highly cathodic to the primary
alloy matrix phase, thereby capable of causing severe localized corrosion. However,
when present in large amounts, these secondary phases can form a continuous network
that can provide a barrier to the corrosion front and retard corrosion, as shown in
Figure 5.2 (Song, 2011).
60
Figure 5.1 Effect of impurities
and alloying elements on
corrosion rate of Mg alloys in
Cd
20
3 wt% NaCl (Makar & Kruger,
1993).
15
10
Corrosion rate (100 mpy)
5
0
6
50
)
–1
40
30
20
Corrosion rate (mm year
10
0
Fe
Ni, Co
Cu
Na, Si, Pb,
Sn, Mn, Al
Alloying element, wt%
Ag
Pb, Sn, Al
Ca
Zn
4
3
21
5
182 Surface Modification of Magnesium and its Alloys for Biomedical Applications
Galvanically accelerating effect
Distribution of a small amount
of discontinuous secondary
phase in a Mg alloy
Corrosion damage of a Mg
alloy with a small amount of
discontinuous secondary phase
Figure 5.2 Schematic representation of dual role of secondary phase of Mg alloys in corrosion
(Song, 2011).
Corrosion barrier effect
Distribution of a large amount
of continuous secondary
phase in a Mg alloy
Corrosion damage of a Mg
alloy with a large amount of
continuous secondary phase
Matrix phase
Secondary
phase
Corroded
area
For example, b-phase in common AZ alloys can either play a role in forming a
corrosion barrier to retard corrosion or a galvanic cathode to accelerate corrosion,
depending on the amount, distribution, and continuity of the b-phase (Figure 5.2).
A fine distribution of b-phase can form a continuous layer and be efficient in providing
the corros ion barrier, whereas the presence of a small amount of discontinuous b-phase
accelerates the microgalvanic corrosion (Song, 2011). In the case of biodegradable Mg
alloys, the microgalvanic corrosion is a serious problem and has been reported in
several previous studies (Bobby Kannan & Singh Raman, 2008; Kirkland et al.,
2010, 2011; Li, Gu, Lou, & Zheng, 2008; Xu, Yu, Zhang, Pan, & Yang, 2007; Xu,
Zhang, Yin, Zeng, & Yang, 2008; Zainal Abidin, Atrens, Martin, & Atrens, 2011).
5.2.2 Localized/pitting corrosion
Mg alloys suffer localized or nonuniform attack as a result of microgalvanic effect (as
described earlier) or due to the electrochemical inhomogeneity at the alloy surface
(Singh Raman, 2004; Singh Raman, Birbilis, & Efthimiadis, 2004; Song & Atrens,
1999). Mg can undergo pitting corrosion at its free corrosion potential in chloride-
containing medium (Tunold, Holtan, Berge, Lasson, & Steen-Hansen, 1977). Besides
the roles of impurities and secondary phases, the uneven distribution of alloying
elements in the matrix can cause development of microgalvanic cells and localized
corrosion.
The localized corrosion and pitting have also been observed during in vitro and in
vivo corrosion studies of Mg alloys for biomedical applications (Witte et al., 2005).
Witte et al. (2005, 2006) reported the pitting corros ion of both LAE442 (Mg—
4 wt% Li—4 wt% Al—2 wt% rare earth) and AZ91D (Mg—9 wt% Al—1 wt% Zn)
Mechanical integrity of magnesium alloys for biomedical applications 183
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Figure 5.3 Localized/pitting corrosion observed in different Mg alloys: (a) AZ61 (Zhou et al.,
2010), (b) AZ91 (Bobby Kannan & Singh Raman, 2008), and (c) Mg-Zn-Mn-Ca (Zhang &
Yang, 2008).
during in vitro and in vivo tests (AZ91D showing a much greater susceptibility to the
localized corrosion/pitting). Several other studies have also reported the susceptibility
of AZ series Mg alloys to localized corrosion/pitting in simulated body fluid (SBF).
Zhou, Shen, and Aung (2010) and Kannan and Raman (2008) observed the localized
corrosion along the residual b-Mg
phase at the grain boundaries and found some
17Al12
pits within the grains (Figure 5.3(a) an d (b)). Localized corrosion and pitting were
observed in the case of some alloys more suitable for implant application, e.g.,
Mg-Zn-Mn-Ca alloys (Zhang & Yang, 2008), which was attributed to the tendency
of Mg
Ca network at the grain boundary to dissolve preferentially during exposure
2
to Hanks’ solution (Figure 5.3(c)). Recently, severe localized corrosion/pitting was
reported for a high-strength Mg alloy, ZX50 (specifically designed for temporary
implant applications in osteosynthesis), during both in vitro and in vivo testing
(Fischerauer et al., 2012; H€anzi, Gerber, Schinhammer, L€offler, & Uggowitzer,
2010; Kraus et al., 2012).
5.2.3 Stress corrosion cracking
A considerable fraction of research on Mg alloys has focused on improving their
general mechanical properties, creep resistance, and corrosion properties. The aspect
that has received relatively less attention is their resistance to stress corrosion cracki ng
184 Surface Modification of Magnesium and its Alloys for Biomedical Applications
(SCC). SCC is a dangerous, complicated, and insidious form of corrosion-assisted failure, and it can cause catastrophic and unexpected fracture of structural components
(Jones, 1992). It is a subcritical crack propagation in a brittle manner due to combined
and interdependent effects of corrosion and tensile stress. SCC involves an optimum
interaction of three parameters: (1) mechanical loading providing tensile stress, (2) a
susceptible alloy, and (3) an environment causing corrosion at the required rate
(Parkins, 1980; Sieradzki & Newman, 1985; Winzer et al., 2005). Consequently,
SCC is relatively less frequent, although failures can be very costly and destructive
when they occur. Unlike other types of corrosion, in the case of SCC, the surface
may appear virtually unattacked, and only a few fine cracks (invisible to naked eye)
initiate and propagate, leading to the brittle failures (Winzer et al., 2005). The most
fundamental and detrimental feature of SCC is that a ductile material that would
have undergone considerable elongation before fracture may suffer embrittlement in
the presence of the corrosive environment.
Mg and its alloys are also reported to be suscep tible to SCC, fracturing at stresses as
low as 50% of the yield stress in common service environments (Winzer et al., 2008;
Winzer et al., 2008a, 2008b). It is estimated that approximately 10e60 Mg alloy
components in aerospace applications alone suffered SCC failures each year during
1960 and 1970 (Winzer et al., 2005). It is also suggested that the occurrence of
SCC may be on increase because Mg alloys are now increasingly used in structural
and automotive applications under load-bearing and harsh environmental conditions.
Increased SCC incidences may also be expected because of increased stresses, as a
result of smaller section sizes of the components in a bid to decrease weight.
5.3 SCC of magnesium alloys: a biomedical perspective
The biomaterials based on Mg alloys suffer from the problems of excessive corrosion
rate and also concomitant hydrogen evolution (Ma & Xu, 2009; Staiger et al., 2006;
Zberg, Uggowitzer, & Loffler, 2009). Mechanical integrity of implanted structures in
the physiological environment is of fundamental importance in effective fracture
fixations and cardiovascular surgeries. Metallic implant devices exposed to corrosive
physiological environment are often also subjected to complex loadings during
service. For example, a permanent hip implant can experience a load up to approximately four times of the body weight during a normal walk, whereas a temporary
cardiovascular stent is continuously subjected to a cyclic loading due to heart beats
(Gu et al., 2010; McCormack, Walsh, Wilson, & Prendergast, 1998). For an adult,
loads in the spine during certain activities may exceed 3500 N (Kirkland, 2012).
The synergistic presence of mechanical loading along with the corrosive environment
may pose complication of sudden fracture of implants due to the phenomenon of
SCC, which can occur even at stresses considerably below yield and design stresses
(Jones, 1992; Singh Raman, 2005), and lead to a premature brittle failure. These brittle cracks generally initiat e at locations of sharp contours such as root of a corrosion
pit. Pre-existing macroscopic/microscopic flaws (e.g., micro-cracks, sharp corners,
and protrusions in implant devices) can also markedly increase the probability of
Mechanical integrity of magnesium alloys for biomedical applications 185
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premature SCC failures and can be locations of SCC crack initiation. Moreover, the
load-bearing implant devices are becoming relatively complex in design, thus also
increasing the probability of high stress intensity points. Furthermore, galvanic, localized, or intergranular corrosion, and even thinning by uniform corrosion, could
initiate SCC in biomaterials (Winzer et al., 2005). In this respect, it is advantageous
to use simple shapes, with an aim of avoiding localized areas of high stress
intensities. Bolted and riveted joints can produce a point of high stress concentrations. It may not always be not possible to avoid having such locations; however,
it may still be possible to reduce stress by some surface treatment such as
electropolishing.
Stress corrosion cracks may propagate undetected to a sudden catastrophic failure
of implant devices during service, which may have many serious consequences such as
painful irritation or inflammation of surrounding tissues, and troublesome removal of
failed devices. In the past, several incidents involving SCC and corrosion fatigue of
traditional implant devices of stainless steels and titanium alloys have been reported
(Bombara & Cavallini, 1977; Bundy, Marek, & Hochman, 1983; Bundy, Vogelbaum,
& Desai, 1986; Hughes & Jordan, 1972; Sivakumar, Kamachi Mudali, & Rajeswari,
1994; Sivakumar & Rajeswari, 1992; Yokoyama, Ichikawa, Murakami, Miyamoto,
& Asaoka, 2002). In this context, it is also important to note that the Mg alloys are
reported to be susceptible to SCC in aqueous environments, including distilled water
and chloride solutions (Winzer et al., 2005). Hence, the characterization of SCC of
biomedical Mg alloys in physiological conditions is a necessity before their actual
service use. However, a very limited attention has been paid on the evaluation of
mechanical integrity/SCC of Mg alloys in body fluid (Bobby Kannan & Singh Raman,
2008; Bobby Kannan, Singh Raman, Witte, Blawert, & Dietzel, 2011; Choudhary &
Singh Raman, 2012).
5.4 Phenomenology of SCC of magnesium alloys
SCC mechanisms are generally dominated by mechanical or electrochemical processes
and often specific to certain combinations of alloy and environment. Transgranular
stress corrosion cracking (TGSCC) is the common mode of SCC in Mg alloys
(Ben-Hamu, Eliezer, Dietzel, & Shin, 2008; Chen, Ai, Wang, Han, & Ke, 2009;
Winzer et al., 2005, 2008a, 2008b), but instances of intergranular stress corrosion
cracking (IGSCC) has also been reported (Bobby Kann an, Dietzel, Blawert, Atrens,
& Lyon, 2008). The fundamental understanding of TGSCC and IGSCC can be found
elsewhere (Winzer et al., 2008, 2008a). In the case of Mg alloys, it is generally
accepted that the SCC processes usually involve hydrogen embrittlement (HE), with
the source of hydrogen being cathodic reaction of the bare metal (Ebtehaj , Hardie,
& Parkins, 1988; Stampella, Procter, & Ashworth, 1984), but there is a little consensus
on a unified SCC mechanism. This section briefly discusses possible mechanisms of
SCC of Mg alloys, influence of alloying elements on SCC, techniques for characterization of SCC in common service environments, and in SBF.
186 Surface Modification of Magnesium and its Alloys for Biomedical Applications
(a) (b)
σ
Mº → Mn+ + ne
σ
Figure 5.4 (a) Continuous crack propagation by dissolution following film rupture, and
(b) cleavage-type fracture ahead of embrittled zone (Winzer et al., 2005). s is tensile stress.
–
σ
Embrittled region
σ
Brittle crack
Slip
5.4.1 Mechanism of SCC in magnesium alloys
SCC of Mg and its alloys has been generally attributed to (1) continuous crack propagation by dissolution at the crack tip (Figure 5.4(a)) or (2) cleavage-type mechanism
by discontinuous crack propagation (Figure 5.4(b))(Winzer et al., 2005).
5.4.1.1 Dissolution mechanisms
The IGSCC observed in Mg-Al alloys is often attributed to the preferential dissolution
of the metal matrix adjacent to the secondary phase precipitates (Mg
alloy). Therefore, a network of grain boundary precipitates facilitates continuous crack
propagation (Winzer et al., 2005). Microgalvanic corrosion causing IGSCC was also
reported by Kannan et al. (Bobby Kannan et al., 2008) for an Al-free rare earth containing alloy, ZE41 (Figure 5.5).
17Al12
in Mg-Al
(a) (b)
Figure 5.5 Fracture surfaces of ZE41 in: (a) 0.5 wt% NaCl—predominant intergranular and
isolated transgranular (arrows) cracking, and (b) distilled water—corrosion along the grain
boundaries (Bobby Kannan et al., 2008).
Mechanical integrity of magnesium alloys for biomedical applications 187
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A film rupture model for SCC of Mg alloys has also been proposed (Winzer et al.,
2005). The localized plastic deformation at the crack tip causes the rupture of passive
film (Jones, 1992), exposing the bare metal to corrosive environment that causes a
rapid dissolution, which leads to crack extension or propagation. The subsequent crack
propagation is governed by the competing processes of film rupture an d repassivation,
as suggested also for Mg alloys by Ebtehaj et al. (1988) and Wearmouth, Dean, and
Parkins (1973).
5.4.1.2 Mechanical fracture mechanisms
The limited slip systems in Mg alloys (hexagonal close-packed (HCP) structure) make
them susceptible to cleavage fracture. A discontinuous cleavage results in the TGSCC
of Mg alloys (Winzer et al., 2005). Various researchers (Chakrapani & Pugh, 1975,
1976; Meletis & Hochman, 1984) have reported that TGSCC in Mg alloys is a result
of alternating processes of electrochemical dissolution and discontinuous crack
advancement due to crystallographic constraints of HCP system. TGSCC resulted in
fracture surfaces consisting of flat and parallel facets separated by perpendicular steps,
which is consistent with cleavage mechanism. The matching and interlocking opposite
fracture surfaces confirmed the occurrence of cleavage, which was difficult to explain
by a dissolution model (Chakrapani & Pugh, 1975, 1976).
Mg alloys are reported to evolve considerable amount of hydrogen even at open
circuit potential (OCP) as well as undergo localized corrosion. Hence, combined
effects of hydrogen-assisted stress corrosion cracking (HASCC) and localized dissolution have been suggested to play an important role in deterioration of mechanical
properties of AZ91D in modified simulated body fluid (m-SBF) (Choudhary &
Singh Raman, 2012). In the case of magnesium alloys, an anodic polarization (which
would normally not facilitate hydrogen generation in other alloy systems) is found to
accelerate the SCC because hydrogen is generated even at such potentials due to the
negative difference effect which is exclusive to Mg alloys (Stampella et al., 1984;
Uematsu, Kakiuchi, & Nakajima, 2012; Winzer et al., 2005).
The TGSCC observed in Mg alloys has been widely attributed to mechanism
involving hydrogen (H). Uematsu et al. (2012) showed that SCC of wrought AZ31
magnesium alloy was dominated by hydrogen embrittlement (HE). They reported
higher crack propagation rate with increasing magnitude of cathodic potential that
facilitated generation of hydrogen. Meletis and Hochman (1984) suggested that the
presence of cleavage features at fracture surfaces could also be attributed to HE.
5.4.2 Influence of alloying elements
Most pure metals are immune to SCC; however, pure Mg has been reported to be
susceptible to SCC (Fairman & Bray, 1971; Lynch & Trevena, 1988; Meletis &
Hochman, 1984; Stampella et al., 1984). Stampella et al. (1984) performed a range
of slow strain rate testing experiments on pure and commercial Mg in Na
different environmental conditions to establish the mechanism of SCC. Both commercially pure and high-purity Mg were reported to be susceptible to SCC at room
2SO4
under
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