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64 Surface Modification of Magnesium and its Alloys for Biomedical Applications
Figure 2.14 Surface morphology of untreated (a, b) and ion-implanted (titanium (c, d) and
titaniumeoxygen (e, f)) WE43 magnesium alloys after immersion in simulated body fluid
for 3 h. Insets in (a) and (c) are the energy-dispersive X-ray spectroscopy analysis performed at
the white corrosion products marked with a red þ.
Adapted from Zhao, Wu et al. (2013) with permission from Elsevier.
modifications, the surface of Mg alloys treated by PIII always exhibited pitting corrosion, and their ability to provide a long-term corrosion resistance remains open. Based
on the characteristic properties of the PIII-treated Mg alloys, it seems that the window
of opportunity of using this method for surface engineering Mg toward the development of degradable implants is rather limited.

Surface modification of magnesium and its alloys: opportunities and challenges 65
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2.6 Laser surface modification
Depending on the kind of interactions between the laser and the material being
treated, as well as the type of microstructural/compositional changes induced at
the surface, laser-based surface modification can be classified as follows: laser surface hardening, laser surface melting (LSM), laser surface alloying (LSA), laser composite surfacing, laser surface cladding (LSC), and laser shock peening (LSP)
(Dahotre, 1998; Singh & Harimkar, 2012). The role of laser surface engineering of
Mg alloys was reviewed by Singh and Harimkar (2012). The effect of different ty pe s
of laser treatments on improving the corrosion resistance of Mg alloys is addressed
by Dutta Majumdar and Manna (2013, chap. 6). Th e importance of surface modifi-
cation of biomaterials by lasers has been addressed by Dahotre, Paital, Samant,
and Daniel (2010), Narayan and Goering (2011),andBandyopadhyay, Balla, Roy,
and Bose (2011). Pulse d l as er deposition is another useful approach to modify t he
surface of Mg alloys.
2.6.1 Laser surface melting
LSM involves heating materials using a high-power laser source, namely,
neodymium-doped yttriumeAlegarnet or carbon dioxide lasers, to their melting
point followed by rapid solidification of the melted zone. Optimizing the conditions
to avoid oxidation of the metal surface is necessary du ring LSM. The common aims
of LSM are (1) refinement of the surface microstructure; (2) homogenization of
composition; and (3) dissolution of precipitates (Dutta Majumdar & Manna,
2011). LSM has been explored as a surface modification method for many Mg alloys.
Guo, Yue, and Man (2005) report ed that the surface of aWE43Mgalloysubjectedto
LSM was free of corrosion even after 4 h of immersion in 3.5% NaCl, whereas its
untreated counterpart showed significant pitting under similar conditions. Abbas,
Liu, and Skeldon (2005) reported that after LSM, the corrosion rate of AZ31,
AZ61, and WE43 alloys in 5% NaCl with a pH of 10.5 at 20
30%, 66%, and 87%, respectively. The improvement in corrosion resistance
exhibited by these three alloys was due to the refinement of a-Mg grains and uniform
redistribution of the b-phase after LSM. Among them, the ability of the AZ61 Mg
alloy to offer a better corrosion resistance than the AZ31 Mg alloy was primarily
due to the higher percentage of Al in it. The s upe rior corrosion resistance off ered
by the WE43 Mg alloy was due to the presence of a rare earth element on the surface.
In addition, in the case of the WE43 Mg alloy, the entire melted layer was not
consumed during the t es t period.
Guan, Zhou, and Zheng (2009) studied the corrosion behavior of LSM of an
AZ91D Mg alloy in SBF. The general and pitting corrosion resistance of the lasertreated AZ91D Mg alloy surface was significantly enhanced by the refined continuous
network of b-Mg
phases and the increased Al concentration in the laser-melted
17Al12
zone. Banerjee, Singh Raman, Durandet, and Mcadam (2011) observed a marginal
improvement in the resistance of an ZE41 Mg alloy to corros ion in 0.001 M NaCl
only during the initial periods of immersion (1 h), which is nullified with a further
C was reduced by about

66 Surface Modification of Magnesium and its Alloys for Biomedical Applications
increase in time. The main reason for the lack of improvement in corrosion resistance
of the LSM ZE41 Mg alloy was because of the formation of surface cracks following
the high thermal stress created by the high input energy of the laser irradiation. This
was substantiated by the smaller extent of surface cracking at a laser power of 820
rather than 1856 W. Another major issue in LSM is the change in microstructural
features and surface roughness in the overlapping tracks during treatment. Coy et al.
(2010) showed a change in microstructure, increase in surface roughness, and forma-
tion of microcracks in the overlapping tracks, as well as the presence of trapped pores,
during LSM of an AZ91D Mg alloy, which increased with an increase in number of
laser pulses (Figure 2.15). Guan et al. (2009), Guan, Zhou, Zheng, and Li (2010),
Figure 2.15 Secondary electron (aec) and backscattered (def) scanning electron micrographs
showing the surface appearance and cross-sectional microstructure of an AZ91D magnesium
alloy after laser treatment using 10 (a, d), 25 (b, e), and 50 pulses (c, f).
Adapted from Coy et al. (2010) with permission from Elsevier.

Surface modification of magnesium and its alloys: opportunities and challenges 67
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and Guan, Zhou, Li, and Zheng (2013a) reported that the formation of a coarse structure in the overlapped area caused by scanning speed provided preferential sites for
pitting corrosion in SBF. According to Guan, Zhou, Zheng, and Li (2013), an increase
in laser power from 40 to 70 W led to the formation of cauliflower-like clusters with
protruding particles on the surface of an AZ31B Mg alloy, which increased its average
surface roughness from 0.11 to 2.69 mm. Trapping of light on the rough surface has led
to a darkening effect of the treated alloy. LSM has been explored as a pretreatment for
PEO of an AZ91D Mg alloy (Wang, Zhou, Liang, & Chen, 2012). The refinement of
microstructure, enrichment of Al, and redistribution of the b-phase (Mg
17Al12
) on the
AZ91D Mg alloy matrix after LSM has enabled the formation of a more compact PEO
coating with fewer cracks on the surface and a higher amount of MgAl
phase in the
2O4
resultant coating. The corrosion resistance of the LSMePEO-treated Mg alloy was
found to be much better than those without the LSM pretreatment. Hence, it is evident
that LSM could offer an improvement in corrosion resistance of Mg alloys in chloridecontaining environments, including SBF, provided necessary precautions are taken
during treatment to avoid the formation of a coarse microstructure in the overlapping
areas and cracks due to the high thermal stress induced by the laser irradiation. In addition, LSM can be used as a pretreatment for Mg alloys for chemical conversion treatments and PEO to improve the reactivity of the alloy, to facilitate a homogeneous
reaction at the surface, and to promote uniformity of the coating.
2.6.2 Laser surface alloying
LSA invol ves adding the preferred alloying elements to the melt pool either by direct
injection or by a preplaced coating followed by laser irradiation of the surface to melt
the alloying elements (injected or preplaced) and the surface of the base material,
generating an enriched melt pool that subsequently solidifies to form an alloyed surface layer. Very high cooling rates (up to 1011 K/s) during resolidification enabled
the formation of a refined microstructure in the alloyed surface. The depth of the
alloyed layer and extent of alloying depends on laser processing parameters such as
laser power, scan speed, laser spot size, and injection rate of alloying elements
(or thickness of the preplaced layer) (Dahotre, 1998). The laser processing parameters
need to be carefully optimized to ensure complete melting of alloying elements and the
formation of strong metallurgical bonding between the alloyed layer and the substrate
(Singh & Harimkar, 2012). It has been reported that an increase in laser power
decreases the average alloying element co ntent in the alloyed zone due to increased
dilution. Also, an increase in scan speed decreases the average composition distribution due to the lower power input at a low interaction time (Dutta Majumdar & Manna,
2013, chap. 6). Galun and Mordike (1996) showed that LSA of Mg with Al and nickel
(Ni) offers better resistance against corrosion. According to Ming, Da, and Chang
(2008), LSA of an AZ91D Mg alloy with AleSi improved its corrosion resistance
in 3.5% NaCl. Paital et al. (2012) reported that LSA of an AZ31B Mg alloy with
Al enabled an improvement in corrosion resistance. In spite of the ability of LSA to
improve the corrosion resistance of Mg alloys in chloride-containing environments,
a proper choice of alloying elements and laser processing conditions is essential to

68 Surface Modification of Magnesium and its Alloys for Biomedical Applications
ensure a defect-free alloyed zone (Dutta Majumdar, & Manna, 2013, chap. 6). In addition, while choosing LSA of Mg alloys for use in biomedical applications, it is important to consider the toxicity of the alloying elements in the human body.
2.6.3 LSC and laser composite surfacing
LSC involves melting the material to be clad on the substrate without significant dilution from the substrate. The rapid solidification of the cladding material led to grain
refinement, extension of solid solubility, and formation of nonequilibrium phases
in the clad layer without significant changes in the microstructure of the substrate
(Subramanian, Sircar, & Mazumder, 1991). In laser composite surfacing, hard ceramic
particles are introduced in the laser-melted surface of the substrate to form a composite
layer after subsequent resolidification. LSC of an MgeZr alloy onto Mg and its alloys
showed improved resistance against corrosion (Su bramanian et al., 1991). LSC of an
AS41 Mg alloy with an AleSi eutectic alloy and its composite with C-short fibers
increased the corrosion resi stance of the Mg alloy by 100 times (Neubert, Bakkar,
& Huang, 2006). In addition, a graded multilayer coating of Ni/copper (Cu)/Al with
good metallurgical bonding at the interfaces between the MgeAl, Ale Cu, and CueNi
layers was succes sfully deposited on Mg by LSC. The LSC-treated Mg specimen
exhibited a large span of passivity of over 500 mV, in contrast to the untreated Mg
substrate, which showed no signs of passivation.
2.6.4 Laser shock peening (LSP)
LSP involves rapid irradiation of the component’s surface with a very high power laser
(power density w10
generation of shock waves (due to expansion of the volume of the plasma plume
formed on the surface) and subsequent alteration of the microstructure/state of stress.
The substrate being treated is usually covered with a thin insulating or absorbing layer
(paint, metal, or plastic tape) and a transparent material (water). Upon laser irradiation,
the generation of plasma induces shock waves, causing plastic deformation at the surface and introducing compressive residual stresses. In addition, the introduction of a
high density of dislocations (Banas, Elsayed-Ali, Lawrence, & Rigsbee, 1990; Banas,
Lawrence, Rigsbee, & Elsayed-Ali, 1990) and the formati on of other phases or twins
(Chu, Rigsbee, Banas, & Elsayed-Ali, 1999) during the process helped to increase the
hardness and yield strength of metallic materials. Moreover, the presence of compressive residual stresses in the material led to a remarkabl e improvement in the fatigue
strength and fatigue life of metallic materials. Because the laser spot size can be
tailored suitably, the LSP can be successfully used for treating components of complex
geometry, and this attribute would be beneficial in treating metallic implants such as
stent materials.
The corrosion behavior of laser peened materials has been addressed by many
researchers. Zhang, You et al. (2010) reported that an AZ31B Mg alloy subjected to
LSP effectively reduced the initiation of stress corrosion cracking (SCC) of the alloy
in 1 wt% NaOH for 500 h at room temperature. Jiang, Zhou, Fan, Huang, and Zhao
9
e1012W/m2) applied at a short pulse (1e50 ns), resulting in the

Surface modification of magnesium and its alloys: opportunities and challenges 69
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(2009) confirmed an improvement in the fatigue life of AZ31B and ZK60 Mg alloys
after LSP. Sealy and Guo (2010) studied LSP of a biodegradable MgeCa alloy to
understand the change in surface integrity and process mechanics and developed a
three-dimensional semi-infinite simulation to predict the topography and residual
stress fields produced by sequential peening. Guo, Sealy, and Guo (2012) reported
that LSP induced unique surface topographies, high compressive residual stresses,
and extended strain hardening, which helps to control the corrosion rate of MgeCa
alloys by more than 100-fold in SBF. According to them, with LSP, the subsurface
of an MgeCa alloy could be strain hardened up to 500 mm, and the maximum
compressive residual stress of the peened surface in the planar direc tion might have
contributed to the decrease in the corrosion rate of the MgeCa alloy. The ability of
LSP to induce compr essive residual stress, to improve fatigue life, and to reduce the
susceptibility of Mg alloys to SCC will be useful in the development of Mg-based
degradable load-bearing implants.
2.6.5 Laser-assisted coloring/darkening and surface patterning
Guan, Zhou, Li, and Zheng (2013b, 2014) studied the effect of femtosecond and
krypton fluoride excimer laser irradiation on the surface evolution of an AZ31 Mg
alloy, resulting in coloration and darkening of its surface following changes in surface
profile and composition. According to them, by adjusting laser parameters according to
the thermal properties of materials and the environmental conditions (oxygen or argon
atmosphere), this type of laser irradiation offers potential applications for the development of Mg-based bio-optical implant devices. Foroozmehr, Alemohammad, Wang,
Toyserkani, and Esmaeili (2011) studied patterning of Mg with silver (Ag) nanopar-
ticles printed by a laser-assisted maskless microdeposition process. The laser-assisted
maskless microdeposition process involves two steps: (1) microdeposition, in which
the deposited material in the form of suspended nanoparticles is atomized to an aerosol
and injected on a moving substrate in a layer-by-layer fashion; and (2) laser postprocessing, in which the laser beam is scanned over the deposited tracks of nanoparticles
for sintering. The Ag film has a relatively uniform thickness with a root mean square
roughness of 48 nm. It is well known that surface patterning of biocompatible materials
could promote cell adhesion, infiltration, and proliferation. Being a bioactive material,
Ag provides better biocompatibility and it is well known for its antibacterial effects.
However, the formation of cracks on the Ag films during sintering is considered a major
limitation in terms of its wide applicability for biomedical applications.
2.7 Other methods of surface modification
2.7.1 Cold spray deposition
Cold spray (CS) coating is a viable method to engineer the surface of Mg and its alloys.
The deposition of CS coatings involves the ballistic impingement of particles, usually
in the size range of 1e100 mm, that are accelerated by a high-velocity stream of gas on

70 Surface Modification of Magnesium and its Alloys for Biomedical Applications
the substrate surface. Upon impact with the substrate, the particles undergo severe
plastic deformation, resulting in flattened splats, which is also accompanied by peening of the substrate. As the temperature of the expanded gas stream that exits the spray
nozzle is relatively lower than the conventional thermal spray processes, it is called a
“cold spray” process. Hence, it is particularly suitable for coatings and substrate materials that are sensitive to heat or oxidation and that would transform or react under the
conditions of a conventional thermal spray process (Champagne, 2007). The performance of CS coatings is largely a function of bond strength between the coating
and the substrate and the residual stress, porosity, and density of the coating. To
achieve a better bond strength, it is important that the mean velocity at which the particles are accelerated exceeds the minimum critical velocity so that the particles have
sufficient kinetic energy to generate extensive plastic deformation of the substrate and
to promote interlocking of the spray splats (Champagne, 2007). Because the particles
are sprayed at relatively lower temperatures, there is little temperature-driven dimensional change, and hence the resultant coatings usually are in a state of compressive
residual stress, which is beneficial to prevent delamination of the coating. The porosity
of CS coatings is generally low, on the order of 1e 5%, and most of them are identified
on the top surface. The following factors, either individually or in combination, are
considered responsible for the occurrence of porosity in CS coatings: (1) rebound of
particles with a velocity lower than the critical velocity; (2) large particles that do not
undergo enough plastic deformation upon impact, generating porosity at the particlee
particle boundaries; (3) the absence of a significant peening effect on the particles
that are already deposited, particularly at the top of the coating; and (4) the sudden
incompatibility in terms of localized heating, stresses, and plastic deformation (Balani,
Laha, Agarwal, Karthikeyan, & Munroe, 2005; Bu, Yandouzi, Lu, & Jodoin, 2011).
Most of the studies of CS coatings on Mg and its alloys concentrate on the deposition of Al and AleAl
composite coatings, particularly to understand the micro-
2O3
structural and interfacial characteristics, to improve the bond strength and density, to
promote formation of intermetallic compounds by heat treatment after coating, and to
improve the corrosion behavior (Bu et al., 2011; Spencer, Fabijanic, & Zhang, 2009;
Spencer, Luzin, Matthews, & Zhang, 2012; Spencer & Zhang, 2009; Tao et al., 2010;
Wang, Qiu, Xiong, Birbilis, & Zhang, 2014; Wang, Spencer, Birbilis, & Zhang, 2010).
Zhang et al. address the role of CS coatings on Mg in Chapter 14 of Volume 2 of this
book. Noorakma, Zuhailawati, Aishvarya, and Dhindaw (2013) recently studied the
deposition of HA on an AZ51 Mg alloy by a modified CS process in which the Mg
alloy was preheated to 400
C for 1 h and the HA powder (average particle size,
4 mm) was sprayed using a high-pressure (10 bar) air nozzle spray at room temperature. According to them, this modification helped to retain the characteristics of HA,
which would otherwise undergo a phase change at high temperature. In addition,
the combination of a higher impact velocity of HA particles and preheating of the
AZ51 Mg alloy enabled better bonding of the HA particles with the Mg alloy and
the formation of a uniform coating. The average thickness and modulus of the resultant
HA coating was 25 mm and 9 GPa, respectively. Immersion in SBF for up to 14 days
showed that the HA-coated AZ51 Mg alloy is bioactive and facilitated apatite
formation.

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Being a low-temperature process, CS is particularly suitable for the deposition of
bioactive coatings such as HAP on Mg and its alloys, in which it would be possible to
limit the extent of oxidation of the substrate as well as the phase transformation of
HAP. As long as the powder particles a re accelerated at a speed that exceeds the minimum critical velocity, the bonding strength between the substrate and the coating
would be good. Peening of the substrate as well as the deposited coatings is likely
to induce compressive residual stress, which would be beneficial in improving the
fatigue strength of the substrate as well as in preventing delamination of the
coating e both attributes are critical in biomedical implant applications. Porosity
of the CS coating is a matter of concern (Figure 2.16). Fortunately, none of these
pores is continuous from the outermost surface of the coating down to the substrate,
suggesting that the coating is impermeable to solution (Tao et al., 2010), which is an
important feature for considering them for magnesium and its alloys in the development of degradable implants. Adopting after spraying a heat treatment at temperatures
closer to the melting point of the substrate for a longer treatment time would offer the
benefits of increasing the bond strength and relaxing the compressive residual stress
(Spencer et al., 2009; Spencer & Z hang, 2009), both of which increase the adherence
of the coating and prevent its delamination. This methodology has many positive
attributes for surface modification of Mg and its alloys toward the development of
(a) (b)
(c) (d)
Figure 2.16 Backscattered scanning electron microscopy images of unreinforced aluminum
(Al) and AleAl
prepared by cold spray deposition: (a) Unreinforced Al coating cold sprayed at 65
(b) unreinforced Al coating cold sprayed at 125
cold sprayed at 65
Adapted from Wang, Spencer et al. (2010) with permission from Elsevier.
composite coatings showing the presence of porosity in the coatings
2O3
C; and (d) Ale50 vol% Al2O3composite coating cold sprayed at 65C.
C; (c) Ale25 vol% Al2O3composite coating
C;

72 Surface Modification of Magnesium and its Alloys for Biomedical Applications
degradable implants. Being a line-of-sight process, the difficulty in coating complex
shapes and the internal surfaces of implants is a major limitation of the CS process
(Champagne, 2007).
2.8 Summary and concluding remarks
This chapter provides a detailed outline of the various surface modification methods
available to modify the surface of Mg alloys and discusses the opportunities and
limitations of each in the development of Mg-based degradable biomaterials. The
salient features are summarized below.
• Imparting self-passivation of Mg alloys through selective oxidation of alloying elements
such as scandium and yttrium is indeed an effective method to control the rate of degradation
of Mg alloys. This approach will be useful when the implant geometry restricts the use of any
coatings. However, the alloying elements should not exhibit any significant toxicity in the
human body.
• Microstructural modification of AZ-type Mg alloys by developing a continuous b-phase
network on the alloy’s surface could also be used to impart a self-passivation mechanism
after the corrosion process. The formation of galvanic cells and an increase in the corrosion
rate are believed to promote deposition of Mg(OH)
and Mg
the porous nature of Mg(OH)
(PO4)2as the primary and secondary corrosion products, respectively. However,
3
would allow permeation of the body fluid and limit the
2
long-term ability of the Mg alloy to protect against corrosion.
• Hydrothermal treatment of Mg alloys enables the formation of an Mg(OH)
hydrophilic and adherent and offers a higher resistance to corrosion than their untreated
counterparts. However, the ability of the Mg(OH)
for the Mg alloys against corrosion in SBF is a matter of concern.
• AHT has the ability to reduce the rate of degradation of Mg alloys. The dense and compact
nature and the amount and crystallinity of MgO are the decisive factors in controlling the rate
of degradation of Mg alloys in SBF.
• Passivation in 1 M NaOH or in m-SBF has limited potential in the surface modification of
Mg and its alloys in the development of degradable implants because of the insufficient
thickness and compactness of the passive films formed in them.
• The difficulty imposed by the intrinsic heterogeneity of Mg alloys, slow reaction kinetics,
and the defective nature of the surface conversion layer formed by the reaction between
the Mg alloys and ILs suggest that there remains much to be explored in generating robust
passivating films using ILs.
• The formation of well-ordered and closely packed SAMs with a high homogeneity, good
order, and better chemical stability is indeed promising. However, the rapid decrease in
the contact angle of some of the SAM-modified Mg alloy surfaces in aqueous solutions
within a short duration of time has raised questions on their long-term chemical stability
in the human body.
• The difference in reactivity imposed by the intrinsic heterogeneity of Mg alloys, the formation of a porous coating with a typical “mud crack” pattern, possibly due to the release of
hydrogen gas and/or dehydration of the coating after deposition, limits the extent of protection against corrosion offered by conversion coatings. Among them, fluoride conversion
coating is promising in terms of its biocompatibility. However, modification of the existing
and Ca3(PO4)2(or Ca10(PO4)6(OH)2)
2
coating, which is
2
coating to provide long-term protection
2

Surface modification of magnesium and its alloys: opportunities and challenges 73
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methodologies to prepare MgF2coatings with desirable characteristics and with a better
understanding of the mechanism of deposition and the influence of volume fraction of
Mg(OH)
on the resistance of the resultant coatings against corrosion needs to be
2-xFx
explored for its potential utility.
• The susceptibility of the sol-gel-deposited BGC coatings for cracking after immersion in
SBF is the major drawback that limits their acceptance as a suitable surface coating of Mg
alloys in the development of degradable implants.
• Silane coating offers only a short-term protection for Mg alloys against corrosion. Alkalinization at the interface during corrosion promotes swelling and decomposition of the SiO
network, while accumulation of corrosion products promotes delamination and blistering
of the coating, thus exposing the Mg alloy to degrade further. However, silane coatings
will be useful for functionalizing compounds with biofunctional properties. Hence,
increasing the bond strength and hydrolytic stability of silane coatings becomes critical in
achieving reasonably good protection of Mg alloys against corrosion.
• It is possible to coat or bind bioactive compounds such as gelatin or albumin. However, the
inability of the precoatings to form a completely dense layer on Mg alloys and the detachment/delamination of the precoating and/or bioactive molecule coating is a matter of concern
in realizing their potential utility.
• Degradable polymer coatings deposited by both dip and spin coating methods did not offer a
long-term corrosion protection. Bulk erosion/degradation leads to detachment of coatings,
resulting in the formation of many micro-/macropores and cracks in the early stage of degradation. Surface-eroding polymer coatings exhibit homogeneous surface erosion from exterior to interior, and this type of coating is beneficial for the development of fully
biodegradable cardiovascular stents.
• The ability to load anti-inflammatory drugs in conducting polymer coatings such as PEDOT
and its subsequent release by electrical stimulation confirmed its potential application in
degradable implant materials with drug delivery applications. However, the inability of
PEDOT coatings to completely prevent the Mg from corrosion, its detachment after multiple
electrical stimulations of drug release, and the mode of degradation of PEDOT under in vivo
conditions are the major concerns.
• Anodizing is useful in developing a passive oxide layer on Mg alloys. However, the adverse
effect on fatigue properties caused by oxidation-induced surface tensile stress, structural
defects in the oxide layer, and substrate “age softening” following the heat associated
with oxide film formation restrict their widespread acceptance. MAO is considered an option
to reduce this risk of fatigue failure. However, because of the higher pore density on the
surface of the MAO coatings of Mg alloys, the effective surface area and the tendency of
the corrosive medium to adsorb and concentrate into these pores are increased, thus limiting
the protection against corrosion offered by them.
• ED is a useful method for depositing HA, calcium-deficient HA, and fluoride-, strontium-,
Zn-, and Si-doped HA coatings. Because Mg alloys have a lower melting point, the difficulty
associated with heat treatment at high temperatures to improve the adhesion of the coatings is
a major limitation in EPD. However, MAO coatings could be used as a pretreatment for
deposition of a variety of inorganic materials with a better adhesion by EPD.
• The major limitation of the coatings deposited by PVD is the presence of pores and pinholes,
which in most cases are difficult to eliminate completely. The formation of a transition layer
to buffer the stress developed by the mismatch between the Mg alloy and the coatings, and an
improvement in deposition methodology by implementing filtered cathodic arc deposition,
could not completely prevent penetration of the electrolyte or sustain their protective properties after long-term exposure in SBF.
2
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