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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5568_Библиотеки_им_академика_М_И_Перельмана
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Plasma electrolytic oxidation/micro-arc oxidation of magnesium and its alloys 209
formed by AC current. However, in DC processing, zirconia particle incorporation in
PEO coatings can also reveal phase transition. The original monoclinic zirconia was
also present in tetragonal modifications when the PEO process was performed in the
same silicate-based electrolyte. In a phosphate-based electrolyte under the same processing conditions, the phase transformation and the reactive incorporation to Mg
2Zr5O12
occurred (Arrabal, Matykina, et al., 2008a).
This generally demonstrates how complex the conditions and interactions of the
various parameters during PEO processing are and how difficult it is to predict the
outcome of the treatment in the sense of coating composition.
To introduce antibacterial/antimicrobial properties, Necula et al. (2009) produced
MgO-Ag nanocomposite coatings on Mg alloys by PEO treatment in alkaline
silicate-based solutions containing Ag-nanoparticles. EDS analysis showed that the
surface content of Ag as well as the roughness of the coating increased with increasing
concentration of Ag in the electrolyte. The GDOES (glow discharge optical emission
spectroscopy) elemental profile showed that Ag was incorporated throughout the
coating and Ag enrichment occurred mainly at the coating/substrate interface. The
morphology of the coating also changed with varying concentrations of Ag nanoparticles in the electrolyte. The surface density of pores seems to be reduced with
increasing concentration of Ag in the electrolyte. However, a rough and fragile coating
was formed with higher concentration of Ag nanoparticles.
Calcium phosphate is another interesting phase for biomedical applications.
Seyfoori, Mirdamadi, Seyedraoufi, Khavandi, and Aliofkhazraei (2013) have demon-
strated that a nanocomposite PEO coating containing biphasic calcium phosphate can
be obtained if 10 g/L hydroxyapatite nanopowder is added to the treatment electrolyte
(7 g/L Na
SiO3þ 5 g/L NaF and 2 g/L KOH). Incorporation of these nanoparticles
2
could alter the microstructure of the coatings by blocking the structural pores of the
MAO-derived films; in this way, it could enhance significantly the corrosion resistance
and apatite-forming ability of the pure MAO film.
To enhance the corrosion resistance of the coating, the sealing of pores is quite
crucial. A study by Song et al. (2012) showed that PEO coating formed in alkaline
aluminate electrolyte containing increasing concentration of titania nanoparticles
forms denser and thicker PEO coating, which showed improved corrosion protection.
However, PEO coating formed in alkaline silicate-fluoride electrolyte with glycerol
showed a slightly decreased growth rate with increasing concentration of titania
(TiO
) sol (Wang et al., 2009). The surface morphology of the coating show ed that
2
the size of discharge channels decreased with the addition of titania sol in the electrolyte. The x-ray diffraction (XRD) pattern of latter study showed that peaks corresponding to Mg
SiO4became broader with addition of titania sol in the electrolyte,
2
indicating formation of amorphous phases in the coating. However, the coating formed
with the highest concentration of titania showed the worst corrosion protection in
simulated body fluid. The presence of the amorphous phase in the coating was
made responsible for the lowest corrosion protection.
Blawert et al. (2012) showed that clay particles can be added to the electrolyte and
significant sealing of the pores can be achieved. It was also discussed that the formation of coating material not only occurs at the coating/substrate interface but also at the
210 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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coating/electrolyte interface. However, incorporation of clay particles in the PEO
coating significantly turned crystalline phases into amorphous ones. The presence of
amorphous phases in the coatings accelerated degradation of the coatings (Lu, submit-
ted for publication).
From these observations, it can be concluded that a dense PEO coating with crystalline microstructure might be more promising for durable corrosion protection of
Mg-based biomedical implants. The positive effect of amorphous phases can result
in a more uniform but faster degradation. However, one should also keep in mind
that Mg alloys are considered to be degradable implants, and therefore a completely
protective coating should not be developed. If the coating is not soluble, it can have
negative effect s on the human body if the coating fractures and insoluble coating particles remain after the substrate is dissolved.
8.5 Duplex treatments
PEO coatings are generally porous, and pores resulting from the discharge channels
often connect the coating/substrate interface with the outside. Such porous coatings
can cause early degradation of the PEO-coated Mg implants. Therefore, pores must
be sealed to enhance the durability of PEO-coated Mg implants in corrosive environments; in addition, the sealants should also be biomaterials. Commonly used sealants
are TiO
applied by either dip-coati ng or electrodeposition processes. There are several reports
of duplex treatment of PEO-coated Mg metal/alloys.
alkaline phosphate, alkaline silicate or silicate-based sol-gel did not cover all the
discharge pores. However, the corrosion resistanc e of the sealed PEO coatings was
improved (Malayoglu, Tekin, & Shrestha, 2010). Shi, Ng, Wong, and Cheng (2009)
fabricated a composite coating on a pure Mg metal for improving its corrosion resistance in Hank’s solution. A thick and porous oxide layer was first formed by PEO in a
silicate/fluoride containing electrolyte, and then a top TiO
on the porous layer by sol-gel dip-coating followed by hydrothermal treatment.
Figure 8.7 shows that all the discharge pores are sealed after TiO
thermal treatment. However, several cracks developed after sealing. The formation of
cracks was suggested to be due to the stress generated by filling the pores. Both electrochemical impedance spectroscopic and anodic polarization measurements showed
an increase in the initial corrosion resistance by about 30 times due to the composite
coating. Immersion tests also showed that the coated samples were more stable over
time, and the degree of corrosion damage was much reduced compared to that of
bare Mg. However, the coating failed in 10 days, which is not enough protection for
practical purposes.
They employed PEO treatment of Mg alloy followed by sol-gel TiO
sealing, samples were annealed to 250
coating before and after sealing and subseq uent mechanical scratching. No cracks
, hydroxyapatite (HA), phosphate, silicate and chitosan, which are often
2
An attempt at sealing the open pores of PEO coating on Mg alloys by immersion in
sealing layer was formed
2
sealing and hydro-
2
Zhang, Bai, et al. (2012) devised a scheme to avoid cracking of the sealant layer.
sealing. After
C. Figure 8.8 shows the surface of the PEO
2
Plasma electrolytic oxidation/micro-arc oxidation of magnesium and its alloys 211
Figure 8.7 SEM micrograph showing surface morphology of PEO-TiO2.
Reprinted from Shi et al. (2009) with permission from Elsevier.
were observed due to sealing (right image). Furthermore, it was observed that filled
discharge pores were tightly filled. Such sealed discharge pores enhanced the corrosion
resistance in simulated body fluid significantly. The corrosion current of the sealed
PEO coatings decreased remarkably by about three orders of magnitude compared
to the bare alloy, and the coating remained intact for a longer time in corrosive simulated body fluid.
Like TiO
, HA is another highly biocompatible and bioactive material, which can
2
also be incorporated in PEO coatings by duplex treatment. A PEO coating was formed
on the Mg-Zn-Ca alloy and a layer of HA was deposited on the coating by electrodeposition (Gao, Guan, et al., 2011). The HA layer effectively filled all the defects, such
as pores and cracks; furthermore, no cracks developed after sealing, as shown in
Figure 8.9. EDS mapping showed the presence of Ca in the outer layer, suggesting
that the outer layer is HA. In addition, scanning electron microscopy (SEM) showed
Figure 8.8 SEM image PEO coating before (left) and after (right) sealing.
Adapted from Zhang, Bai, et al. (2012) with permission from Elsevier.
212 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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(a)
(b)
(c)
Figure 8.9 (a) Surface morphology and (b) cross-section morphology of composite coatings on
Mg-Zn-Ca alloy. (c) EDS mapping showing the presence of Ca in the outer layer, suggesting that
the outer layer is HA.
Reprinted from Gao, Guan, et al. (2011) with permission from Elsevier.
Plasma electrolytic oxidation/micro-arc oxidation of magnesium and its alloys 213
that HA appears in the form of nanorods. This duplex coating remarkably decreased
the corrosion current by about three orders of magnitude and ennobled the corrosion
potential by more than 100 mV in SBF. Immersion in SBF also promoted the formation of nanoflakes of HA from nanorods, revealing a phenomenon of biomineralization
by precipitation of calcium phosphates. This phenomenon was reported only for HA
nanorods.
A similar study by Sreekanth & Rameshbabu (2012) also showed the improvement
of biodegradation resistance in PEO-coated AZ31 Mg alloy by the electrodeposition of
HA. However, HA can also be derived from calcium phosphate coating, even by immersion in SBF. Liu, Hu, Ding, and Wang (2011) formed a PEO coating on pure Mg
samples and a calcium phosphate layer was formed by a chemical method (calcification). X-ray crystallography revealed that the duplex coating was composed of a
mixture of HA, dicalcium phosphate dehydrate (DCPD) and MgO. In a later study,
Liu et al. reported the influence of the deposition temperature. At low temperatures,
the previously mentioned mixture is formed, whereas at higher temperature (67
C)
mainly the formation of HA is observed (Liu, Tang, Li, & Hu, 2014).
After immersion of the duplex-coated Mg in SBF for 147 h, large flake-like structures of dicalcium phosphate dihydrate possibly dissolved in SBF and the fine porous
structure was changed to needle-like structures of HA (Liu et al., 2011). Longer immersion times in SBF confirmed the already observed change of the coating. The
authors called the change from the original HA/DCPD coating to bonelike apatite
‘self-adjustment’, which should result in better corrosion resistance as well (Liu
et al., 2014). A PEO coating containing dicalcium phosphate dihydrate was also
formed when PEO treatment was carried out in an electrolyte containing EDTA-Ca
with subsequent hydrothermal treatment of the coating at low temperature. However,
high-temperature hydrothermal treatment led to filling of the discharge channels and
densification of the coating accompanied with disappearance of the dihydrate. Potentiodynamic polarization measurements of the hydrothermally treated coating showed a
remarkable decrease in anodic current in Hank’s solution over a wide range of polarized potential (Chang, Tian, Liu, & Duan, 2013). In contrast, Zhang, Ma, Chen, and
Wei (2013) have produced Ca-P coatings on Mg-Zn-Zr alloy using a three-step
process, consisting of anodizing, pre-phosphatizing (5 h, 37
C, saturated Na2HPO
solution) and deposition of Ca-P coating (up to 48 h, 70C, solution of Na2HPO
$12H2O þ Ca(NO3)2$4H2O). Compared with direct deposition, anodizing as the
4
pretreatment followed by pre-phosphatizing makes the CaeP coating denser and the
binding force with the Mg alloy increases remarkably, while the corrosion potential
shifts to higher values.
As demon strated by Lin et al., posttreatment with HF (dipping for 48 h, 30
C,
48 wt% HF) can be another option to seal or modify an MAO coating on a biodegradable magnesium alloy (Lin, Tan, Wan, et al., 2013). After the treatment, the porous
structure of the MAO coating disappeared, and magnesium fluoride becam e the
main component of the coating. The wear resistance was not affected much. A slightly
reduced corrosion resistance was observed, but the initial alkalization effect of the
MAO coating was well controlled; thus, no haemolysis took place for the
HF-treated MAO coating.
4
214 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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Another option to increase the corrosion resistance and reduce the degradation rate
is the combination of PEO coatings with polymer coatings. The combination of a PEO
coating and a polyether imide (PEI) top-coat improved the corrosion resistance
in 0.1 M NaCl solution (electrochemical impedance spectroscopy, EIS) and salt spray
(5% NaCl) remarkably (Srinivasan, Scharnagl, Blawert, & Dietzel, 2010). However,
corrosion tests in physiological solutions have not been performed yet. An interesting
feature for biomedical applications is the controlled and adjustable degradation behaviour of this PEI polymer over a wide range (Scharnagl & Blawert, 2013b); thus,
biodegradation should be still available for the whole coating system.
Chitosan can be also used for successful sealing of MAO coatings on Mg-Zn-Ca
alloys. The sealing is obtained by a 5-min dipping of the MAO specimen in a chitosan
solution (1% chitosan dissolved in 1 wt% acetic acid) and subsequent drying at room
temperature. The duplex treatment reduces the corrosion current densi ty in SBF by
nearly three orders of magnitude compared to the uncoated alloy (Bai et al., 2012).
An example for the combination of MAO and multilayers composed out of subsequent
layers of chitosan and polystyrene sulfonate was given by Liu, Pan, Yang, Cai, and
Chen (2012). With duplex treatment, the corrosion resistance of the WE43 alloy in
SBF could be significantly increased.
8.6 Performance of PEO coatings in biomedical
applications
In actual human physiological environments, orthopaedic implants and cardiovascular
stents are under mechanical stress and are subject to corrosion and wear. Even if the main
mechanical properties are generally provided by the substrate, the biomedical coatings
should have adapted mechanical properties; for example, they should be able to follow
the strain of the magnesium substrate without failure, protect the substrate for a certain
time and also be biodegradable. The PEO coatings’ morphology and composition can be
controlled over a wide range; thus, the properties can be fine-tuned to offer reasonable
mechanical strength, wear and corrosion resistance. Therefore, they have a high potential for biomedical applications. However, characterization of PEO coatings in in-vivo
environments is still limited. This section gives a general overview of recent studies
of in vitro and in vivo performances of PEO-coated magnesium alloys.
8.6.1 Mechanical properties
Mg alloys have excellent mechanical properties for application as implants in loadbearing parts of the body, such as the human tibia. The fracture toughness of Mg alloys
is higher than that of ceramic biomaterials, and the elastic modulus and compressive
yield strength of magnesium are closer to those of natural bone than other metallic implants. Thus, Mg alloys help to reduce or avoid the ‘stress shielding effect ’— a facto r
that reduces the stimulation of new bone growth (Staiger et al., 2006). To endure
the stress during movement of bones, PEO-coated Mg alloys should have similar
Plasma electrolytic oxidation/micro-arc oxidation of magnesium and its alloys 215
mechanical properties to that of substrate alloys. PEO coatings are formed due to the
conversion of metallic substrates into oxide, so there is strong adhesion of the coating
with substrate. The average compr essive stress in human tibia is about
4.047 0.217 MPa (Fukubayashi & Kurosawa, 1980; Wang Fei, 2006). PEO coatings are hard enough to resist compressive stress. However, degradation properties
may be different under stress. Wang, Gao, Wang, Zhu, and Guan (2012) studied the
biodegradation of a PEO-coated cast Mg alloy under compressive stress in Kokubo’s
SBF and found that the applied compressive stress changed the degradation mode of
coated samples. The applied stress caused certain peeling-off locations on the coating,
and continuous increase of pH value demonstrated faster degradation of the coating
under stress. However, precipitation of Ca-P salts inside the micropores and on the surface of coated samples bearing stress was still observed, suggesting that the coating
still had protective and osteoconductive abilities.
The effect of a PEO coating on the strength retention of in vivo and in vitro degraded
magnesium implants (WE43) was studied by Imwinkelried, Beck, Iizuka, and Schaller
(2013). The difference in strength retention of coated and noncoated implants was most
pronounced after 12 weeks and vanished with longer implantation times. The authors
found that approximately 80% of the strength was retained at 12 weeks of implantation,
which should allow a stable fracture fixation during the period typically needed for
fracture healing. Another study reported by Xia, Zhang, Lu, and Geng (2013) came
to similar conclusions. In spite of the loss of corrosion resistance and loss of strength
of PEO-coated Mge4.0Zne0.2Ca alloy after 30 days of immersion in SBF, its mechanical integrity was still judged to be enough for bone fixtures.
However, under bending load conditions, PEO coatings are also exposed to tensile
stresses. Coatings may crack if they cannot follow the strain of the substrate. However,
SCC (stress corrosion cracking) experiments with PEO-coated Mg alloys in very mild
chloride containing test solutions (ASTM D1384 solution) have demonstrated that the
coatings may crack, but they do not severely flake off and they do improve the SCC
resistance compared to the uncoated alloy (Bala Srinivasan, Blawert, & Dietzel, 2008;
Srinivasan, Blawert, Dietzel, & Kainer, 2008).
If the coatings are exposed to alternating loads, the fatigue properties are important.
Unfortunately, there is no informat ion available on the performance in physiological
solutions. The performance of PEO-coated magnesium alloys in air and in technological chloride solutions suggests that the performance is similar compared to uncoated
material as long as the coatings are not too thick (<20 mm). For more information, see
the review by Blawert et al. (2006).
PEO coatings are normally composed of hard crystalline oxides. Therefore, PEO
coatings are highly resistant to mechanical wear. The wear properties of PEO coatings
are also controlled by the microstructure of the coatings. There are several reports
about the characterization of mechanical (wear) properties and microstructure of
PEO coatings formed on Mg alloys for industrial and biomedical applications. The
hardness of PEO-coated Mg alloys is controlled by crystalline phases present in the
coatings, which in turn can be controlled by processing parameters, electrolytes and
substrate composition. Liang et al. (2005b) fabricated PEO coatings in alkaline silicate
electrolyte with and without KF; they found that the microhardness of the coating is
g
y
216 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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Figure 8.10 Microhardness of Mg alloy
700
substrate and PEO coatings formed in
different solutions.
Reprinted from Liang et al. (2005b) with
permission from Elsevier.
600
500
400
300
200
Microhardness (Hv)
100
0
Untreated
M
allo
Coating formed
in solution 1
Coating formed
in solution 2
higher for the PEO coating formed with KF addition (Figure 8.10). The higher microhardness may be due to the formation of dense microstructure and crystalline MgF
phases in the coating. The wear rate for the coating formed in the electrolyte containing
fluoride is lower, and the reason should be the formation of additional crystalline
phases (Figure 8.11). Wear resistance of PEO coating formed in silicate electrolyte
is higher than that formed in phosphate electrolyte, which is related to the higher hardness of the silicate-based coating (Rapheal, Kumar, Blawert, & Dahotre, 2011).
Finally, the knowledge about the influence of PEO coatings on the mechanical
properties of Mg alloys under in vitro and in vivo test conditions is still quite limited.
The tests available suggest that the performance is at least better than the performance
of the uncoated alloys and sufficient for certain applications, such as fracture fixation.
2
Figure 8.11 Wear rates of Mg
4.0
alloy substrate and PEO coatings
formed in different solutions.
Reprinted from Liang et al.
(2005b) with permission from
Elsevier.
/ Nm)
3
mm
–4
0.20
0.15
0.10
Wear rate, (× 10
0.05
0.00
3.5
3.0
2.5
Untreated
Mg alloy
Coating formed
in solution 1
Coating formed
in solution 2
Plasma electrolytic oxidation/micro-arc oxidation of magnesium and its alloys 217
However, alternating loads and fatigue properties have not been studied to the knowledge of the authors and are still a concern.
8.6.2 Corrosion/biodegradation
Mg is a highly electropositive metal and is readily corroded in aqueous medium containing chloride ions. In body fluids, there is a considerable amount of chloride ions
and magnesium is easily corroded. In an early test of magnesium as a biomaterial,
Lambotte (1932) reported that Mg corroded too rapidly in vivo, leading to complete
disintegration within 8 days and producing a significant amount of gas beneath the
skin. Such a fast disintegration of Mg led to failure of the fracture healing and a significant amount of gas liberated at the detached implant/tissue interface, which further
hindered fracture healing. PEO coatings have the potential to control the degradation
rate of the substrate. Corrosion of PEO-coated Mg alloys is controlled by the microstructure and composition of the coating. There are several reports about the biodegradation of PEO-coated Mg alloys.
Zhang, Zhao, Wu, Wang, and Wu (2007) reported that the degradation and
hydrogen evolution in Hank’s solution was significantly suppressed by PEO coating
of AZ91D alloys. Electrochemical polarization showed that open-circuit potential
shifted significantly to the positive direction by PEO coating, showing significant
passivation by the coating (Figure 8.12). The corrosion current of the alloy was
dramatically decreased by five orders of magnitude due to the PEO coating. However,
the PEO coating degraded just after immersing the coated sample for 1 day in Hank’s
solution because the coating was porous. Lin, Tan, Zhang, et al. (2013) produced PEO
coatings at different voltages of 230, 300, 370 and 450 V on ZK60 Mg alloys in
–0.1
–0.2
–0.3
–0.4
–0.5
E (volts)
–0.6
–1.5
–1.6
–1.7
1E-10 1E-9
MAO sample
1E-8 1E-7 1E-6
Untreated sample
1E-5 1E-4 1E-3 0.01 0.1 1
2
)
I (A / cm
Figure 8.12 Polarization curves for PEO-treated sample and untreated one (potential measured
against saturated Hg-Hg
2Cl2
electrode).
Reprinted from Zhang et al. (2007) with kind permission from Springer Science and Business
Media.
218 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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alkaline silicate-fluoride electrolyte. PEO coatings with different microstructures were
identified, but only the coating formed at 450 V was highly dense and uniform and
showed the best corrosion protection. However, after immersion of this coating for
1 month in Hank’s solution, several voids appeared in the coating, and formation of
these voids might be due to dissolution of MgO.
Therefore, further improvement in PEO coating composition and microstructure is
required to decrease the degradation rate of PEO-coated Mg alloys. Most of the PEO
coatings contain MgO as a major phase, and MgO is unstable in aqueous medium.
Therefore, replacing or reducing the content of MgO would be a strategy to slow
down the degradation if required. This can be achieved by modifying the electrolyte
for PEO, producing other phases that are supposed to be more biocompatible.
PEO coating containing CaSiO
different concentrations of Ca(H
was formed in alkaline silicate electrolyte with
3
2PO4)2
at different constant voltage modes. The
pore density and morphology of coatings largely changed with varying concentrations
of phosphate as well as the magnitude of constant voltage, as shown in Figure 8.13.
Increasing the concentration of the phosphate is effective in enhancing the coating
growth rate and reducing the pore density. The coating formed in the electrolyte containing 15 g/l phosphate at 500 V showed the lowest biodegradation over 4 weeks of
immersion in SBF (Wang et al., 2013). Tang and Wang (2013) reported that a CaTiO
based PEO coating can be formed on AZ31 magnesium alloys by using an electrolyte
consisting of sodium hydroxide NaOH (6 g/L), potassium fluorotitanate K
and calcium glycerophosphate C
tive and biocompatible CaTiO
3
PCa (6 g/L). The PEO coating containing bioac-
3H7O6
significantly suppressed the biodegradation rate of the
TiF6(4 g/L)
2
alloy substrate. Additionally, a uniform layer containing spherical particles with a
high potential to form apatite was observed on the coating surface after biocorrosion
(7 days of exposure to SBF). It was reported that the formation of this layer could
enhance the biocompatibility of the coated Mg alloy.
An electrochemical corrosion study (pot.-dyn. polarization and EIS) of the performance of a pulsed constant current silicate-based PEO coating on pure Mg in simulated body fluid revealed the much better corrosion performance compared to the
uncoated specimen (Alabbasi, Bobby Kannan, Walter, St€ormer, & Blawert, 2013).
The coating is composed out of a thin dense inner layer, which offers the main corrosion resistance, and a thick porous outer layer. Postdegradation analysis and similar
breakdown potential observed in the potentiodynamic polarization curves of the uncoated and the PEO-coated magnesium suggest that the porous outer layer has stabilized the inner compact layer, most likely by inhibiting the free flow of aggressive ions
towards the inner layer by promoting precipitation (Mg
OH and Ca3(PO4)2) on the
2PO4
porous layer surface.
However, the performance depends very much on the open porosity. Sealing is
required if the lifetime of PEO coatings in physiological solutions should be extended
and the dissolutions should be retarded or slowed down. Sealing treatments can always
be considered as duplex treatments, and possible improvements have already been discussed. However, one should keep in mind that the various sealing techniques available have to be carefully selected for the intended application to obtain the best
performance. Chu et al. have tested the degradation behaviour in SBF and simulated
-
3
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