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34 Surface Modification of Magnesium and its Alloys for Biomedical Applications
different types of alkaline solutions, namely, Na2HPO4,Na2CO3, and NaHCO3for
24 h followed by heat treatment at 773 K for 12 h. During the initial alkaline treatment,
depending on the type of alkaline solution, compounds such as magnesium carbonate,
calcium carbonate, magnesium sodium phosphate, and calcium phosphate were deposited on the surface. The se deposits were loose and defective and accelerated the corrosion rate of the MgeCa alloy. However, during the subsequent heat treatment, these
loose surface layers were fused and consolidated, which was accompanied by the
oxidation of the subst rate, resulting in the formation of a compact MgO layer on the
surface of the MgeCa alloy (Figure 2.4). The average thickness of the MgO layer
formed after AHT in Na
HPO4,Na2CO3, and NaHCO3are 13, 9, and 26 mm, respec-
2
tively. After AHT, the surface roughness of the MgeCa alloy increased. AHT in all the
three solutions improved the corrosion resistance of the MgeCa alloys; based on the
extent of protection against corrosion, they were ranked as follows:
AHT-NaHCO
> AHT-Na2HPO4> AHT-Na2CO3. The increase in surface area of
3
the sample after AHT has enabled the formation of calcium phosphate. The cytotoxicity evaluation showed that AHT of the MgeCa alloy did not induce toxicity to L929 cells.
Nan, Yangmei, and Bangcheng (2013) soaked commercially pure Mg (99.95%
purity) in 5 mL of 1 M NaHCO
at 60C for 24 h followed by heat treatment in air
3
at 773 K for 10 h. According to them, the AHT reduced the rate of Mg degradation,
supported by the slow change in the pH of SBF and smaller volume of hydrogen gas
released during corrosion. After immersio n in SBF for 24 h, the AHT led to the
formation of Ca
Ca
is one of the intermediate products of the biomineralization process, which
2P2O7
. According to El Kady, Mohamed, and El-Bassyouni (2009),
2P2O7
has the potential to regulate the onset of calcification and could promote mineralization in the body. The AHT Mg samples pretrea ted in serum cell culture medium for
24 h showed a smaller extent of adhesion and spreading of MG63 cells after 4 and
6 h of cell culture but exhibited more spreading with sphere-like characteristics after
8 h of cell culture.
Figure 2.4 Scanning electron micrographs taken at the surface (a) and at the cross-section (b) of
a magnesiumecalcium alloy (1.4 wt% calcium) after alkaline treatment in sodium bicarbonate
for 24 h followed by subsequent heat treatment at 773 K for 12 h.
Adapted from Gu, Zheng, Cheng et al. (2009) with permission from Elsevier.
Surface modification of magnesium and its alloys: opportunities and challenges 35
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The findings of Li et al. (2004), Gu, Zheng, Cheng et al. (2009), and Nan et al.
(2013) clearly show that AHT decreased the rate of degradation of Mg and its alloys
in SBF, and the formation of an MgO layer is responsible for the observed improvement in resistance against corrosion. Deposition of MgO on the surface of Mg and its
alloys could be accomplished by heat treatment. However, the amount and crystallinity
of MgO formed exclusively by heat treatment were smaller than those formed by AHT.
The MgO layers prepared only by heat treatment are porous and nonprotective,
whereas those prepared by AHT are relatively dense and compact. The dense and
compact nature and the amount and crystallinity of MgO are decisive factors in
controlling the rate of degradation of Mg and its alloys in SBF.
2.3 Chemical surface modifications
2.3.1 Chemical passivation
Passivation of the surface of Mg is the simplest approach in surface modification.
Passivation of commercially pure Mg (99.9% purity) by soaking in 1 M NaOH for
24 h or in modified SBF (m-SBF) at 37
(2009) as a possible means of surface modification of Mg. The passive layer formed
on Mg after immersion in 1 M NaOH was very thin (on the order of a few nanometers)
and composed of MgO/Mg(OH)
2
m-SBF was relatively thicker, on the order of tens of micrometers, and composed of
amorphous carbonated calcium and magnesium phosphate. Both types of passive
films, however, were not highly protective in m-SBF, which was attributed to the
insufficient thickness and compactness of the passive films formed in 1 M NaOH
and the higher porosity and roughness of the passive film formed in m-SBF. Both types
of passive films showed a significantly larger number of human HeLa and mouse
fibroblasts cells on their surfaces than those observed on the polished Mg surface.
The passive film formed on Mg in 1 M NaOH showed a significant increase in the
cell survival rate. The passive film formed on Mg in m-SBF seems to promote a large
number of cell nuclei during initial cell adhesion, whereas it does not support cell
spreading or favor the formation of actin cytoskeleton after 24 h of cell culture. The
inability of the passive film formed on Mg in m-SBF to control the rate of corrosion
resulted in a large increase in pH that led to cell death. Hence, it is evident that
chemical passivation using NaOH and m-SBF has limited potential in the surface
modification of Mg and its alloys toward the development of degradable implants.
C for 5 days was explored by Lorenz et al.
. The passive film formed on Mg after immersion in
2.3.2 Reaction with ionic liquids
Ionic liquids (ILs) are salts that exist as liquid at temperatures below 100C, and they
are usually composed of a combination of organic cations and organic/inorganic
anions. The interaction between ILs such as trihexyl(tetradecyl)phosphonium
diphenylphosphate ([P6,6,6,14][DPP]), trihexyl(tetradecyl)phosphonium bis(trifluoromethanesulfonyl) amide ([P66614][NTf2]) , and trihexyl(tetradecyl)phosphonium bis
36 Surface Modification of Magnesium and its Alloys for Biomedical Applications
(2,4,4-trimethylpentyl)phosphinate ([P66614][(iC8)2PO2]) and Mg alloys, namely,
ZE41 and AZ31, was studied by various groups (Efthimiadis et al., 2010; Forsyth
et al., 2009; Howlett et al., 2010; Huang, P. et al., 2013; Latham , Howlett, MacFarlane,
& Forsyth, 2011). The findings of their studies indicate that the interaction between ILs
and Mg alloys is not a simple physisorption of the anions of the ILs on the surface of
Mg alloys; instead, it most likely involves a chemical or electrochemical process. The
nonreversible nature of the cyclic voltammetric scans of AZ31 Mg alloy in ([P66614]
[(iC
) further confirms this attribute (Latham et al., 2011). The surface films
8)2PO2
formed on Mg alloys have a thickness of a few tens of nanometers, in which the inner
layer is dominated by metal oxides/hydroxides while the outer layer mainly consists of
organics, most likely entrapped breakdo wn products and components of ILs. These
surface films offered reasonable protection for the Mg alloys against corrosion in a
chloride-containing environment. However, there are many challenges to overcome
before realizing the potential utility of ILs for protection of Mg alloys against corrosion. The intrinsic heterogeneity of Mg alloys poses a major challenge in generating
a uniform surface film. Acid pickling pretreatment and application of a potential
bias during treatment with ILs are some possible means of achieving better surface
homogeneity of Mg alloys, whereas the addition of suitable amounts of water in ILs
enhances its reactivity (Murase et al., 2013). In spite of these modifications, the
slow reaction kinetics warrant a longer treatment time (24 h in most cases), and the
resultant surface films are by no means free of defects. The curren t status of the role
of ILs for corrosion protection of Mg alloys recently was reviewed by Huang et al.
(2013). According to them, much research still is needed to understand the possible
means of generating robust passivating films on reactive metal surfaces such as Mg
in the presence of ILs. Until then, the possibility of using ILs for controlling the
rate of corrosion of Mg alloys toward the development of degradable implants is
limited.
2.3.3 Self-assembled monolayers
The spontaneous formation of a highly ordered monolayer structure of certain organic
compounds, particularly those with a long alkyl chain, on the surface of various substrates is referred as a self-assembled monolayer (SAM). The head groups of these
organic compounds anchor to the surface, while their hydrophobic tail prevents the
approach of water molecules toward the surface. The interfacial properties of the
SAM are largely determined by the type of the head group and the length of the tail
groups. SAMs form by a simple chemisorption process, making them strongly adhere
to the surface. The ability of SAMs to establish conformal contact regardless of the
shape of the substrates and the capability to control their thickness and composition
offer a wide window of opportunity to explore their potential applications in many
areas. The kinetics and mechanism of formation of SAMs was reviewed by Schwartz
(2001). Liu, Yu, Zhou, and Wu (2006) successfully formed SAMs of sodium dodec-
anoate (C
(C
18H35O2
12H23O2
Na) on an AZ91D Mg alloy in ethanol and showed that these SAMs could
offer a corrosion protection efficiency of up to 98.5% in 0.2 M NaCl. A distinct trend
Na), sodium palmitate (C16H31O2Na), and sodium stearate
Surface modification of magnesium and its alloys: opportunities and challenges 37
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observed in the three SAMs is that the longer the alkyl chains of alkylcarboxylate, the
better the extent of corrosion protection. This is because of the increase in the thickness
of the SAMs with an increase in the length of the alkyl chains of these compounds. An
increase in treat ment time of up to 24 h also showed an increase in corrosion protection
efficiency, probably because of a better alignment and packing of the SAMs. The schematic structure of the n-alkylcarboxylate monolayer on the surface of the Mg alloy
with C
carboxylate is depicted in Figure 2.5. Gao, Qiao, and Xin (2010) formed
18
SAMs of stearic acid (SA) on heat-treated (773 K for 10 h) 4N-Mg. According to
them, a combination of heat treatment and SAMs decreased the rate of degradation
of Mg from 0.11 to 0.05 mm/year and induced calcium phosphate deposition in
SBF. The improvement in corrosion resistance following this treatment also was evidenced by the marginal decrease in the diameter of the Mg sample by 0.2 mm after
12 weeks of implantation in rabbits when compared with the uncoated sample, which
exhibited a larger decrease in diameter of 0.93 mm. In addition, when the Mg samples
modified by the combination of heat treatment and SAMs were implanted into the
rabbits, no infections occurred, and the Mg
2þ
concentration in blood after 2, 6, and
12 weeks remained in the natural range.
Ishizaki, Okido, Masuda, Saito, and Sakamoto (2011) and Ishizaki, Teshima,
Masuda, and Sakamoto (2011) formed SAMs of alkanoic and phosphonic acids on
an AZ31 Mg alloy using liquid phase and vapor phase methods. Accordingly, a
phosphonic acid-derived SAM-mod ified AZ31 Mg alloy offered better resistance to
corrosion than the unmodified alkanoic acid-derived alloys. They also reported that
Surface
normal
Oxygen
Carbon
Hydrogen
Film
thickness
Figure 2.5 Schematic structure of the n-alkylcarboxylate monolayer on the surface of a
magnesium alloy with C
CeO bond normal to the substrate surface.
Adapted from Liu et al. (2006) with permission from Elsevier.
18
Extended
molecular
length
θ
carboxylate. The monodentate bonding was estimated assuming a
38 Surface Modification of Magnesium and its Alloys for Biomedical Applications
the adhesion and chemical stability of SAMs of alkyl phosphonic acid prepared by a
simple immersion method was very low when compared with those prepared by the
vapor phase method. These inferences clearly point out that SAMs should be wellordered and closely packed, with their hydrophobic terminal methyl group oriented
outward, and that there exists a strong van der Waals interaction between the alkyl
chains, thus endowing them with a high homogeneity, good order, and better chemical
stability to achieve a better protection of Mg alloys against corrosion. It has been
reported that water contact angles could be a good indicator of the nature of SAMs;
in general, those that exhibit high hydrophobicity offer better protection against corrosion. Salman and Okido (2012) reported that the SAMs of oleic and stearic acids on
AZ31 Mg alloy formed in ethanol showed larger contact angles than those formed
using acetone and hexane, probably because of the formation of magnesium ethoxide
on the Mg alloy’s surface. Hence it is evident that selection of a suitable solvent also is
a key in the formation of well-ordered and homogeneous SAMs.
Grubac et al. (2012) suggested that SAMs of palmitic and stearic acids have much
potential for designing specific interface architectures in degradable implant technology. They formed SAMs of palmitic and stearic acids on bare as well as anodized
AZ91D alloy and evalua ted their corrosion resistance in Hank’s solution at 37
C.
Their findings further confirmed that the longer the hydrocarbon chains of alkylcarboxylate, the better the corrosion protection of the SAMs formed by it. In addition,
they showed that the anodic oxide film provided an excellent base for the chemisorption of carboxylic acids, resulting in a significant decrease in the rate of degradation of
the Mg alloy. Grubac et al. (2013) studied the effect of SAMs of dodecylphosphonic
acid and octadecyl-phosphonic acid, in both as-deposited and heat-treated (120
C for
5.5 and 24 h) conditions, on the corrosion resistance of AZ91D Mg alloy in Hank’s
solution at 37
C. Accordingly, in the as-deposited condition, the intermolecular interaction (Van der Waals and hydrogen bonding) was apparently stronger than the
substrateemolecule interaction, whereas after heat treatment the transformation of
the phosphonic acid monolayer to a phosphonate enabled a strong chemical bond
with the surface, thus leading to a significant increase in protection against corrosion.
Based on the findings of the role of SAMs for protection of Mg alloys against corrosion, it is evident that the formation of well-ordered and closely packed SAMs with a
high homogeneity, good order, and better chemical stability is the key step in
achieving better corrosion protection of Mg alloys toward the development of degradable Mg-based biomaterials. 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 human body fluid.
2.3.4 Chemical conversion coatings
2.3.4.1 Cerium-, nicotinic acid-, titanate-, and sodium
bicarbonate-based conversion coatings
Cui et al. (2011) studied the formation of cerium conversion coatings on an AZ31 Mg
alloy. The cerium conversion coating was dense, composed of CeO
, CeO, Ce2O3,
2
Surface modification of magnesium and its alloys: opportunities and challenges 39
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MgO, Mg(OH)2, and aluminum oxide (Al2O3), effectively reduced the rate of the
AZ31 Mg alloy when compared with the untreated alloy after 24 h of immersion in
Hank’s solution, and showed excellent biocompatibility. Deposition of the nicotinic
acid-based conversion coatings improved the resistance of the AZ31 Mg alloy to
corrosion, and it did not indicate any cytotoxic activity in a normal adult stem cell culture, suggesting its good biocompatibility in cell culture in vitro (Bikulcius,
Rucinskien_e, Bukelskien_e, Selskis, & Grigucevicien_e, 2012). Yang, Tsai, Huang,
and Lin (2012) studied the deposition of a titanate conversion coating on AZ31 Mg
alloy and evaluated the corrosion behavior of the coated Mg alloy in 0.05 M NaCl,
0.1 M sodium sulfate, and after 24 h of the salt spray test (ASTM B117). The titanate
conversion-coated Mg alloy offered a better corrosion resistance than the uncoated
alloy. The effect of deposition of a chemical conversion coating by immersion in saturated aqueous NaHCO
solution on the corrosion resistance of AZ31 and AZ61 Mg
3
alloys in 0.6 M NaCl was studied by Feliu, Samani ego, El-Hadad, and Llorente
(2013). The resultant coating was composed of carbonates of magnesium and sodium
and aluminum hydroxide and offered reasonable protection when compared with its
untreated counterpart. The formation of a typical “mud crack” pattern, possibly caused
by the release of hydrogen gas and/or dehydration of the coating after deposition,
limits the extent of protection against corrosion offered by these coatings.
2.3.4.2 Phosphate conversion coatings
Deposition of phosphate conversion coatings has been explored by many researchers
as a surface modification method to improve the corrosion resistance of Mg and its
alloys (Hu, Meng, Chen, & Wang, 2012; Li, Lian, Niu, & Jiang, 2006; Van Phuong,
Lee, Chang, Kim et al., 2013; Van Phuong, Lee, Chang, & Moon, 2013; Xu, Zhang, &
Yang, 2009). Deposition of a Zn phosphate coating is one of the most well-studied
systems and was recently reviewed by Van Phuong, Lee, Chang, Kim et al. (2013).
It has been reported that the experimental conditions should be optimized carefully
to achieve a good-quality deposit. Being a conversion coating, the resultant phosphate
coating is porous in nature, which limits its fullest potential in improving the corrosion
resistance of Mg alloys (Hu et al., 2012; Li et al., 2006; Van Phuong, Lee, Chang,
Kim et al., 2013; Van Phuong, Lee, Chang, & Moon, 2013). Xu et al. (2009) used
Ca(H
2PO4)2$H2
bath and deposited brushite coating with small amounts of Zn on the surface of an
MgeMneZn alloy. The brushite coating provided good protection to the Mg alloy
against corrosion in SBF. However, in spite of its ability to reduce the corrosion
rate, it fails to completely stop the degradation of Mg in SBF. As a surface coating,
the brushite coating can improve the surface biocompatibility of Mg alloys because
it can be transformed into hydroxyapatite (HA) phase in SBF. A detailed account of
phosphate conversion treatment of Mg alloys for biomedical applications is presented
by Zhang in Chapter 2 of Volume 2 of this book. In addition, the deposition of strontium phosphate and manganese phosphate coatings on Mg alloys and their possible
implications in biomedical applications are addressed by Chen et al. in Chapter 15
of Volume 2 of this book. Deposition of Mg(OH)
O and Zn(H2PO4)2$2H2O as main components of a phosphating
-HA and HA coatings also was
2
40 Surface Modification of Magnesium and its Alloys for Biomedical Applications
explored using the chemical conversion method (Chen, Birbilis, & Abbott, 2011;
Wang, Huang et al., 2013). Deposition of calcium phosphates and HA assumed signif-
icance because of their bioactive properties. More details on the various methods of
deposition and characteristics of calcium phosphates and HA are covered in Chapters
3, 7, and 10 of Volume 2 of this book.
2.3.4.3 Phytic acid-based conversion coatings
Phytic acid-based conversion coatings assumed significance because of their nontoxic
and biocompatible nature as well as the chelating ability of phytic acid with Mg and its
alloys. Jianrui, Yina, and Weidong (2006) studied the deposition of phytic acid-based
conversion coatings on an AZ91D Mg alloy as a function of the concentration of phytic
acid, pH, temperature, and immersion time. Accordingly, phytic acid coatings deposited under optimum conditions (concentration of phytic acid: w0.5e1%;
temperature: w25e60
reasonably good protection of the AZ91D Mg alloy against corrosion in 3.5% NaCl.
Cui et al. (2008) studied the mechanism of formation of phytic acid-based
conversion coatings on AZ91D Mg alloy using 5 g/L of phytic acid at 20
tion of time. Accordingly, in an aqueous solution, the phosphate group of the phytic
acid is ionized, forming phytic acid radicals with a different number of phosphate radicals, phosphate hydrogen radicals, and hydroxyl ions. The phytic acid ion combined
with the surface Mg ions to form a magnesium phytate complex on the surface of
the Mg alloy. The phytic acid conversion coating exhibited a good bonding strength
(5B as per ASTM D 3359-90) and offered reasonable protection against corrosion in
3.5% NaCl. Ye, Zheng, Wang, Xi, and Li (2012) pointed out that phytic acid coatings
deposited on WE43 Mg alloys at pH 3 and 5 were relatively thicker than those coated at
pH 8 and 10; among them, those coated at pH 5 offered a marginal increase in resistance
against corrosion in SBF (Figure 2.6), better cell viability, and an acceptable hemolysis
rate. The findings of Jianrui et al. (2006), Cui et al. (2008), and Ye, Zheng et al. (2012)
clearly indicate that in spite of its complexing ability, nontoxicity, biocompatibility, cell
viability, and hemolytic ability, the presence of cracks is a serious limitation in terms of
achieving higher corrosion resistance with phytic acid-based conversion coatings.
Gupta, Mensah-Darkwa, and Kumar (2013), Gupta, Mensah-Darkwa, Sankar, and
Kumar (2013), and Chen, Huang, and Peng (2013) made some modifications in the
methodology of deposition to overcome this limitation. In one of the approaches, Gupta,
Mensah-Darkwa, Sankar et al. (2013) deposited the phytic acid conversion coating by
immersing Mg in a phytic acid solution (50% water) at room temperature for 3 h, drying
at 60
C for 3 h, followed by immersion in an SA solution (1 mg/mL in ethanol) at
80
C for 2 h, with a thickness of w2.3 mm. The posttreatment in an SA solution filled
the microcracks present in the phytic acid conversion coating, improved the smoothness
of the coating, and reduced the corrosion current density of phytic acid conversioncoated Mg in PBS from 1.16 10
Sankar et al., 2013). In another approach, the phytic acid-coated Mg was subjected
to a heat treatment at 95
(w1.4%) in the film thickness, the heat treatment helped to annihilate the microcracks
C, pH w3e5, and treatment time: w30e60 min) offered a
C as a func-
6
to 1.03 106A/cm2(Gupta, Mensah-Darkwa,
C for 1 h under a vacuum. In spite of a very slight shrinkage
Surface modification of magnesium and its alloys: opportunities and challenges 41
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20
18
2
16
14
12
10
8
6
4
2
Volume of evoled hydrogen / ml / cm
0
–2
PH = 3
PH = 5
PH = 8
PH = 10
Control
0 20 40 60 80 100 120
Immersion time / h
Figure 2.6 Comparison of the volume of hydrogen that evolved during the immersion of an
untreated and phytic acid conversionecoated WE43 magnesium alloy (at various pHs) in
simulated body fluid as a function of immersion time.
Adapted from Ye, Zheng et al. (2012) with permission from Elsevier.
present in the phytic acid conversion-coated layer, with no formation of oxides. The
heat treatment after coating reduced the corrosion current density of phytic acid
conversion-coated Mg in PBS from 1.16 10
6
to 9.27 107A/cm2(Gupta,
Mensah-Darkwa, & Kumar, 2013). Chen, Wan et al. (2013) suggested that pretreatment
of Mg with alkaline in 3 M NaOH at 60
phytic acid layer, with uniform distribution on the surface following the establishment
of a covalent bond via the hydrolysis/neutralization reaction between the OH
and H
O3PO-R (myo-inositol-1,2,3,4,5,6-hexakisphosphate) on the phytic acid mole-
C for about 24 h helped to immobilize the
on Mg
cule. The formation of a dense and homogenous phytic acid conversion coating by this
methodology decreased the rate of degradation of Mg in PBS. The improved bonding of
the phytic acid molecules and the clinically acceptable hemolysis ratio of the coated Mg
suggests that it has potential application in the development of degradable Mg-based
cardiovascular stents.
2.3.4.4 Fluoride conversion coatings
Deposition of fluoride conversion coatings on Mg and its alloys has been studied by
many researchers, particularly because of their suitability for biomedical applications
and their ability to serve as a pretreatment for bioactive coati ngs. A detailed account of
fluoride conversion coatings of Mg alloys is presented by da Conceicao in Chapter 1 of
Volume 2 of this book, and their suitability for biomedical applications is addressed
here. Fluoride coatings offer many advantages.
• Fluoride conversion coatings reduced the corrosion rate of Mg and its alloys under both in
vitro and in vivo conditions (Figure 2.7), confirming their ability to control the localized
increase in pH and accumulation of hydrogen gas (Carboneras, García-Alonso, & Escudero,
42 Surface Modification of Magnesium and its Alloys for Biomedical Applications
Figure 2.7 Surface morphologies of a fluoride-treated (a, c, e, g) and bare (b, d, f, h) AZ31B
magnesium alloy after immersion in simulated blood plasma for different durations: 3 days
(a, b); 7 days (c, d); 14 days (e, f); and 30 days (g, h).
Adapted from Yan et al. (2010) with permission from Elsevier.
Surface modification of magnesium and its alloys: opportunities and challenges 43
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2011; Chiu, Wong, Cheng, & Man, 2007; da Conceicao, Scharnagl, Blawert, Dietzel, &
Kainer, 2010; da Conceicao, Scharnagl, Dietzel, & Kainer, 2012; Drynda et al., 2010;
Drynda, Seibt, Hassel, Bach, & Peuster, 2013; Witte et al., 2010; Yan et al., 2010).
• It offered a good biocompatibility (Carboneras et al., 2011; Drynda et al., 2010, 2013;
Thomann et al., 2010; Witte et al. , 2010), demonstrated good clini cal tolerance in a rabbit
model (Thomann et al., 2010), revealed no adverse effects during surgical implantation
(Drynda et al., 2013), and showed no significant signs of gas accumulation at the implantation site (Drynda et al., 2013), suggesting its better tissue compatibility.
• It ensured a gradual degradation, which helped the Mg alloy to maintain a better mechanical
property for as long as 45 days in SBF, thus meeting one of the important requirements of a
biodegradable material (Pereda et al., 2010; Yan et al., 2010).
• It offered an improved cell response for MG63 cells and MC3T3-E1 cells after 72 h of cell
culture (Li, Li et al., 2013) and promoted the adhesion and proliferation of 3T3 cells (Ye,
Chen, You, & Liu, 2010).
• It enhanced the deposition of calcium phosphate, suggesting its ability to accelerate the rate
of mineral deposition and promote bone healing (Bakhsheshi-Rad, Idris, Abdul-kadir, &
Daroonparvar, 2013; Li, Li et al., 2013).
• It demonstrated a significant increase in bonelike fraction on Mg alloy sponges, as evidenced
by more rapid bone adaptation, better tolerance, higher vascularization, and good osseointegration (Lalk et al., 2013).
• The low solubility of Mg fluoride (MgF
decreased its leaching rate, as evidenced by the absence of an elevated fluoride concentration
in regions adjacent to the bone (Drynda et al., 2010; Witte et al., 2010).
• The fluoride ions released from MgF
Alvarez, & Fernandez Lorenzo de Mele, 2011; Pereda, Alonso, Gamero, del Valle, &
Fernandez Lorenzo de Mele, 2011).
• The MgF
biofilm-forming pathogens: Escherichia coli and Staphylococcus aureus (Lellouche,
Kahana, Elias, Gedanken, & Banin, 2009).
• It exhibited no harm to the red blood cells (Ye et al., 2010), while it could enhance the lyso-
somal activity under specific conditions (Grillo et al., 2011).
• It served as a good pretreatment for the deposition of bioactive coatings such as HA and
dicalcium phosphate dihydrate (DCPD) obtained by electrodeposition (ED) and electrophoretic deposition (EPD) methods (Bakhsheshi-Rad, Idris, & Abdul-kadir, 2013; Jo et al.,
2011; Rojaee, Fathi, & Raeissi, 2013a), and it increased the adhesion and bonding strength
of these coatings (Bakhsheshi-Rad, Idris, & Abdul-kadir, 2013).
coating effectively restricted biofilm formation against two common nosocomial
2
) coating in aqueous physiological solutions(0.13 g/L)
2
coating did not show any cytotoxic effects (Grillo,
2
Based on these characteristics, MgF2-coated Mg and its alloys is considered suitable for cardiovascular stents (Bakhsheshi-Rad, Idris, Abdul-kadir et al., 2013; Drynda
et al., 2010) and for biodegradable implants in endoprothesic applications (Yan et al.,
2010). In spite of these numerous advantages, the acceptance of MgF
alloys as candidate materials for degradable implants still remains limited because
the MgF
coating prepared by chemical conversion method (immersion in
2
20e50 wt% HF at 300 K for 6e168 h) is usually thin (w1e3 mm), porous, and could
offer only a limited increase in the corrosion resistance in SBF (Bakhsheshi-Rad, Idris,
Abdul-kadir et al., 2013; Carboneras et al., 2011; Chiu et al., 2007; da Conceicao et al.,
2010, 2012; Drynda et al., 2010; Pereda et al., 2010, 2011; Thomann et al., 2010;
Witte et al., 2010; Yan et al., 2010; Ye et al., 2010). Chen, Song, Zhang, Li,
-coated Mg
2
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