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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5612_Библиотеки_им_академика_М_И_Перельмана
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Surface design of biodegradable magnesium alloys for biomedical applications 105
3.4.2 Ca-P based coatings
Ca-P coatings such HA and tricalcium phosphate (TCP) are often considered as
osteoconductive materials and have been widely used to construct new bones and
promote osteointegration on biomedical implants because calcium and phosphorus
are t wo major elements in bone tissues. Recently, various techniques such as electrodeposition and hydrothermal methods have been applied to deposit Ca-P coatings on
Mg alloys.
Electro deposition is a promising method for biomedical Mg-based implants
because it can form uniform coatings on porous substrates or implants with a complex
shape at a low deposition temperature. Song, Shan, and Han (2008) used electrodeposition to coat AZ91D Mg alloy with bioactive Ca-P coatings. The electrolyte solution
was prepared from 0.1 M Ca(NO
pH value was adjusted to 4.3. Electrodepositon was carried out at a stable cathodic
potential of 4 V for 2 h at room temperature. The as-deposited coating consisted of
dicalcium phosphate dehydrate (CaHPO
(Ca
(PO4)2, b-TCP) (Figure 3.13(a)).
3
,0.06MNH4H2PO4, and 10 mL/L H2O2and the
3)2
$2H2O, DCPD) and b-tricalcium phosphate
4
(a)
1000
800
DCPD
*
*
o β-TCP
*
600
*
o
o
*
o
*
o
o
0
20
30 40 50 60 70 80 90
HA
HA
HA
0
20 30 40 50 60 70 80 90
HA
2θ / (0)
HA
2θ / (
HA
o
*
HA
0
)
(b)
400
Intensity / cps
200
500
400
300
200
Intensity / cps
100
Figure 3.13 Surface morphology and XRD patterns: (a) As-deposited coating and (b) HA
coating.
Source: Song et al. (2008) with permission from Elsevier.
106 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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Usually, the electrodeposition reactions on the Mg alloy surface are as follows
(Kuo & Yen, 2002; Song et al., 2008):
Stage I: Reduction reaction of H
2
PO
and HPO
4
2
4
2H2PO
2HPO
Stage II: Ca
and Ca
3
Ca
3Ca
þ 2e/2HPO
4
2
þ 2e/2PO
4
2þ
reacting with HPO
(PO4)2(b-TCP), respectively.
2þ
2þ
þ HPO
þ 2PO
2
þ 2H2O/CaHPO4$2H2O (3.4)
4
3
/Ca3ðPO4Þ
4
2
þ H2[ (3.2)
4
3
þ H2[ (3.3)
4
2
4
2
and PO
3
to form CaHPO4$2H2O (DCPD)
4
(3.5)
DCPD and b-TCP are the precursors of HA, which is the stable calcium phosphate
form in an alkaline solution. The as-deposited coating was further immers ed in 1 M
NaOH solution for 2 h at 80
was transformed into a uniform HA (Ca
C to obtain the HA coating. Afterward, the coating
(PO4)6(OH)2) one, as shown in
10
Figure 3.13(b), and the corrosion resistance of AZ91D Mg alloy in the simulated
body fluid (SBF) was increased.
Coatings fabricated by the traditional cathodic electrodeposition process using a
static potential tend to be loose and porous and have low adhesion. One reason is
that a concentration polarization is easily formed because ion diffusion from the solution to the substrate surface is too slow. The other is that H
as a result of reduction of H
O. To overcome these hurdles, Wang, Guan, Wang, Ren,
2
is produced on the cathode
2
and Wang (2010) used pulsed power. The electrolyte was prepared from 0.042 mol/L
Ca(NO
5.0 by diluted HNO
, 0.025 mol/L NH4H2PO4, and 0.1 mol/L NaNO3and the pH was adjusted to
3)2
and (CH2OH)3CNH2. They obtained a soluble Ca-deficient hy-
3
droxyapatite (Ca-def HA) coating on the Mg-Zn-Ca alloy substrate by pulsed eletrodeposition. The Ca/P atomic ratio of the as-deposited coating was about 1.33
(within the range between 1.33 and 1.65). By regulating the pulse amplitude and
width, the Ca-def HA coating had better adhesion to the Mg-Zn-Ca alloy and the
lap shear strength increased to 41.8 2.7 MPa. The ultimate tensile strength and
time of fracture measured from the coated Mg-Zn-Ca alloy were larger than those
from the uncoated one, thus offering benefits in supporting fractured bone healing
for a longer time. In addition, the Ca-def HA coating also improved the corrosion resistance appreciably in SBF.
Chemical solution deposition was developed by Hiromoto and Tomozawa (2011)
to prepare Ca-P coatings on Mg alloys. In their experiment, the solution was prepared
with ethylenediaminetetraacetic acid calcium disodium salt hydrate (C
10H12N2O8
Na2Ca, Ca-EDTA), potassium dihydrogenphosphate (KH2PO4), and sodium hydroxide (NaOH). They first mixed 250 mmol/L Ca-EDTA and 250 mmol/L KH
2PO4
and
then used a NaOH solution to adjust the pH to 8.9. HA coatings were obtained on
-
Surface design of biodegradable magnesium alloys for biomedical applications 107
Figure 3.14 Surface and cross-sectional morphology of the hydroxyapatite (HA) coated AZ31
magnesium alloy.
Source: Hiromoto and Tomozawa (2011) with permission from Elsevier.
AZ31 Mg alloys after treatment at 363 K for 6 h. Those HA coatings possessed a novel
microstructure consisting of an inner dense layer and outer coarse layer as shown in
Figure 3.14. The inner layer on the AZ31 was composed of dome-shaped and densely
packed precipitates, while the outer layer comprised rod-like crystals growing from
each dome in the radial direction. It was also found from the corrosion tests that the
Mg ion release and corrosion current density were remarkably reduced. Hiromoto,
Tomozawa, and Maruyama (2013) studied the fatigue properties of HA-coated
AZ31 Mg alloys. The HA coating prepared by single-step chemical solution deposition consisted of an outer porous HA layer, an inner continuous HA layer, and a
thin intermediate MgO layer. In the tensile test, the HA coating microscopically
showed neither cracks nor detachment at 5% static elongation (1.5% residual strain).
With further elongation under tensile stress, cracks were formed perpendicularly to the
tensile direc tion and fragments of the coating detached from the fracture inside
the inner continuous HA layer. The fatigue strength at 10
7
cycles (fatigue limit) of
the HA-coated and mechanically polished AZ31 was about 80 MPa and 90 MPa,
respectively. The slight decrease in the fatigue limit observed from the HA coating
was attributed to small pits with a depth of about 10 mm. The HA coating remained
on the specimen without cracks after 10
7
cycles at the fatigue limit providing about
3% cyclic elongation.
3.4.3 Polymer-based coatings
Biodegradable polymers such as poly L-lactic acid (PLLA), poly ε-caprolactone (PCL),
and poly glycolic acid (PGA) have been approved for human clinical applications and
become a promising option to improve the initial corrosion resistance and cell compatibility on Mg alloys to meet healing requirements. Dipping, spraying, and spinning are
common preparation methods and some examples are described below.
Wong et al. (2010) prepared porous polymeric membranes on AZ91 Mg alloy.
They first mixed polycaprolactone (PCL) with the average molecular weight (M
about 80,000 g/mol and dichloromethane (DCM) and then deposited the polymerbased membrane layer by layer on the sample surface using a custom-designed
spraying device. In the spraying process, the device was equipped with air-flow and
temperature control, thereby standardizing the thickness, homogeneity, and adhesiveness of the polymer-based membrane. The air-flow pressure and spraying temperature
)of
n
108 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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were 276 kPa and 37C, respectively. The spraying process was conducted at 50%
humidity, 22
C, and atmospheric pressure. The polymeric membranes reduced the
degradation rate while preserving the bulk mechanical properties during degradation.
The polymer-coated samples showed better cytocompatibility with eGFP and SaOS-2
osteoblasts than the uncoated samples and higher volumes of new bone were observed
on the coated samples by microcomputed tomography. Histological analysis indicated
no inflammation, necrosis, and hydrogen gas accumulation during degradation.
Xu and Yamamoto (2012) used spin coating to prepare uniform, nonporous, and
amorphous poly
L-lactic acid (PLLA) and semi-crystalline poly ε-caprolactone
(PCL) films on extruded Mg substrates. Spinning is a typical coating process that involves depositing a polymer solution onto a solid substrate as the substrate was rotated
at a high speed. When the solvent evaporates and the dissolved polymer covers the
substrate, a layer of homogeneous thin polymer film is produced. The experimental details are shown as follows. PLLA with two different molecular weights of 50,000 (low
molecular weight, LMW) and 80,000e100,000 (high molecular weight, HMW) and
PCL with two different molecular weights of 40,000 (LMW) and 70,000e100,000
(HMW) were used. The polymers were weighed to the desired proportions and dissolved in chloroform to obtain 5% (w/v) solutions. Each polymer solution was dropped
onto the Mg sample placed on a spin coater by a micropipette, and the polymer film
was prepared by spin-coating. They investigated the effects of the coating thickness,
adhesion strength between the coatings and substrates, polymer molecular weight,
and different polymers on the corrosion resistance and differentiated cell functions
on the coated Mg-based materials. The PLLA film showed better adhesion strength
to the Mg substrate than the PCL one. For both PLLA and PCL, the LMW films
were thinner and exhibited better adhesion strength than the HMW ones. According
to the pH measurements of the cell culture medium and quantification of released
2þ
Mg
during the cell culture, the corrosion resistance of the Mg substrate was
improved by the polymer films to a different degree. In addition, all the polymeric
films enhanced the cytocompatibility during incubation for 7 days.
3.4.4 Physical vapor deposition coatings
PVD is a modern coating technique commonly used by the industry. Compared to
other methods such as CVD and thermal spraying, the required deposition temperature
in PVD is often lower and so it is very suitable for Mg alloys. Moreover, PVD is
friendly to the environment and is regarded as a green technique that can substitute
for industrial electrochemical plating and anodic oxidation. Last but not least, many
coating species from metals to ceramics can be used. Researchers have used it to
improve the corrosion resistance of Mg and Mg alloys in saline solutions. Hoche
et al. (2005) used sputtering to prepare CrN and Al
cast alloy, and Altun and Sen (2005) conducted DC sputtering to deposit AlN coatings
on AZ31, AZ61, AZ63 and AZ91 Mg alloys. Wu (2007), Wu, Zeng, and Yuan (2008)
and Wu, Ding, Zeng, Wang, and Yao (2009) also used sputtering to coat AZ31 Mg
alloys with Al and Ti metallic coatings. Most of the samples showed improved corrosion resistance in saline solutions.
coatings on AZ91D Mg die
2O3
Surface design of biodegradable magnesium alloys for biomedical applications 109
Xin, Liu, Zhang, Huo, et al. (2008) used cathodic arc-deposition to deposit ZrO
coatings on Mg alloys for biomedical applications because zirconia has good chemical stability and favorable biocompatibility. Commercial extruded AZ91 Mg alloys
were used in their experiment. Argon sputter cleaning was conducted at a bias of
1000 V for about 30 min before deposition. A zirconium transition layer was first
deposited for about half an hour, followed by deposition of the zirconia coating
for about 3 h at a bias of 100 V. Finally, a 1.5 mm-thick ZrO
/Zr bilayered structure
2
with good adhesion was obtained. EIS measurements disclosed that the corrosion
resistance of the coated alloy was significantly improved. Electrolyte penetration
eventually deteriorated the protection of the coating after long exposure in the
SBF. In addition, they also deposited Al
(Xin, Liu, Zhang, Jiang, et al., 2008 )
2O3
and ZrN (Xin et al., 2009) on Mg alloys. Most of these PVD coatings are no nbiodegradable, and so inflammation may occur if they are broken and remain in the human
body for a prolonged period. Synthetic apatites exhibit excellent biological properties, such as biocompatibility, bioactivity, lack of toxicity, or inflammatory and
immunitary responses, and also have relatively high bioresorbability (Jaime,
Michele, José, Stéphanie, & Christophe, 2013). They can also be prepared by phys-
ical vapor deposition such as sputtering (Boyd, Du ffy, M cCann, & Meenan, 2008;
Yamashita, Matsuda, Arashi, & Umegaki, 1998). Thus, it is a good candidate using
PVD on Mg alloy in the future.
3.4.5 Ion implantation
Ion implantation is different from the aforementioned coating techniques. It can
provide the possibility of introducing different species into a substrate independent
of thermodynamic limitations such as solubility. Besides, an ion implanted layer
does not have an abrupt interface, thereby avoiding layer delamination that plagues
coatings. Conventional line-beam metal ion implantation has been attempted to
modify the properties of magnesium substrates. For example, cerium (Ce) ion implantation improves the corrosion resistance of AZ31 Mg alloy (Wang, Zeng, Yao, Wu,
& Lai, 2008) and yttrium (Y) ion implantation enhances the oxidation resistance of
AZ31 magnesium alloy (Wang, Zeng, Wu, Yao, & Lai, 2007).
Zn as one of the vital elements in the human body has been considered and
implanted into pure Mg substrate by conventional beam-line metal ion implantation.
However, after implanting 2.5 10
into pure Mg at 35 kV, the degradation rate increased significantly in SBF. It was
believed to be due to galvanic effects between the metallic Zn-rich surface and Mg
matrix beneath (Wu, Gong, et al., 2011). Al has an electrode potential close to that
of Mg in aqueous solutions, and Al ion implantation can tailor the surface corrosion
resistance of pure Mg. The corrosion resistance in SBF improves significantly and
this enhancement is attributed to the formation of a gradient surface structure involving
a gradual transition from an Al-rich oxide layer to Al-rich metal layer. However, when
the biological properties and toxicity of the alloying elements are considered, the use
of Al is suspected because Al is suspected to be involved with Alzheimer’s disease and
may also cause muscle fiber damage (Wu, Xu, et al., 2012).
17
ions$cm2of Zn using a cathodic arc source
2
110 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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Mg corrosion is affected by the thin surface oxide film and dissolution typically
occurs in the oxide-free areas (Galicia, Pébere, Tribol let, & Vivier, 2009). The native
surface oxide film formed on exposure to air consists of mainly MgO, but MgO is not
stable in an aqueous solution according to thermodynamics and converted to magnesium hydroxide. Cl
in the aqueous solutions can substitute OHforming chloride,
which expedites dissolution of the surface structure (Wu, Feng, et al., 2012). To
improve the chemical stability of the temporary surface on the biodegradable Mg alloy,
some chemically stable phases such as Cr
are needed. Xu et al. (2011) performed
2O3
Cr ion implantation into pure magnesium, but it induced rapid degradation in SBF
similar to Zn ion implantation due to galvanic corrosion. Ensuing oxygen ion implantation produced a thicker oxidized layer composed of chromium oxide that succes sfully retarded surface degradation. Wu, Feng, et al. (2012) applied oxygen ion
implantation to modify the Mg-Nd-Zn-Zr alloy. But unfortunately, no significant
improvement was observed. They performed Cr ion implantation prior to oxygen
ion implantation and attained improved corrosion resistance in SBF due to the form ation of Cr-rich oxide in the surface layer.
Ti and Zr are biologically friendly to the human body and have also been implanted.
Zhao et al. (2013) showed that the surface corrosion resistance on WE43 alloy in SBF
was significantly improved after Ti ion implantation in conjunction with oxygen PIII.
They also conducted Zr and O dual implantation to modify Mg-Ca and Mg-Sr alloys
(Zhao et al., 2014). Besides the improved corrosion resistance in simulat ed physiological environments, the amounts of adherent bacteria on the Zr-O-implanted and
Zr-implanted samples diminished remarkably compared to the unimplanted alloy
(Figure 3.15) and significantly enhanced cell adhesion and proliferation were observed
from the Zr-O-implanted sample. The results suggest that dual zirconium and oxygen
ion implantation is a possible means to avoid inflammation in clinical applications of
biodegradable magnesium alloys.
(a)
(d)
Figure 3.15 Fluorescent microscopic views of magnesium samples after bacteria culturing
for 30 min: (a) Unimplanted Mg-Ca, (b) Zr-implanted Mg-Ca, (c) Zr-O-implanted Mg-Ca,
(d) Unimplanted Mg-Sr, (e) Zr-implanted Mg-Sr, and (f) Zr-O-implanted Mg-Sr.
Source: Zhao et al. (2014) with permission from Elsevier.
(b)
(e)
(c)
(f)
Surface design of biodegradable magnesium alloys for biomedical applications 111
Plasma immersion ion implantation (PIII) can process samples with a complex
shape and is a viable technique to process biomedical artificial joints. Wu et al.
(2014) introduced C
gas into the PIII process and conducted plasma immersion
2H2
ion implantation and deposition (PIII&D) on Mg-Nd-Zn-Zr alloys. PIII&D combines
energetic ion implantation and low-energy plasma deposition, resulting in a thin
diamond-like carbon film formation on the Mg-Nd-Zn-Zr alloy (Figure 3.16). Both
electrochemical and immersion tests reveal enhanced corrosion resistance in the
0.9 wt% NaCl solution. Although the diamond-like carbon film has an excellent barrier
effect against corrosion, defects generated in the deposition process induce corrosion
failure of the plasma-modified Mg-Nd-Zn-Zr alloy in aqueous solutions eventually.
3.4.6 Composite coatings
Composite coatings are attractive as advanced coatings. When two or more constituents are combined to form a layered or mixed structure, the properties of a traditional
coating can be refined to address specific requirements. Various combinations such as
ceramics/ceramics and ceramics/polymers have been proposed.
(a) (b)
t
T
PIII&D
(c) (d)
A
Intensity (a.u.)
0 500 1000 1500
Figure 3.16 (a) Surface appearance of the untreated and treated samples. (b) SEM picture of
surface morphology of the film with the inset showing the magnified surface obtained by AFM.
(c) SEM view of the cross-section of the plasma-modified sample, with the inset showing a
magnified picture of the film. (d) EDS line scan of the film.
Source: Wu et al. (2014) with permission from Elsevier.
Distance (nm)
C
Mg
O
B
2000 2500
112 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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Ca-P coatings not only retard degradation of Mg alloys under physiological conditions, but also have good biocompatibility. However, their fragile nature and structural
heterogeneity induce the loss of integrity possibly reducing the corrosion resistance of
the coated Mg alloys. Wang, Zhao, Chen, Li, and Zhang (2012) used polycaprolactone
(PCL) to preserve the integrity of dicalcium phosphate dihydrate (CaHPO
$2H2O,
4
DCPD) coatings for a longer time because it has good plasticity and uniform structure.
Ca-P coating can avoid direct contact between the PCL and Mg matrix because there is
a negative effect in the interaction in the later degradation stage. Therefore, they fabricated a layered composite coating composed of dicalcium phosphate dihydrate
(DCPD) and polycaprolactone (PCL) on the Mg-Zn alloy as shown in Figure 3.17.
The DCPD coating was synthesized in a 0.042 mol/L Ca(NO
0.025 mol/L NH
4H2PO4
solution by electrodeposition and the DCPD-coated sample
∙4H2Oand
3)2
was immersed in a 2 wt% PCL chloroform solution before drying in air. Compared
to the DCPD-coated alloy, the DCPD/PCL-coated alloy had higher corrosion
resistance as manifested by the elevated corrosion potential, reduced corrosion current
density, and smaller amount of released hydrogen.
Chitosan is a natural biopolymer that exhibits various biological activity including
excellent biocompatibility, biodegradabi lity, osteoconductivity, and antimicrobial
properties. Hahn et al. (2011) prepared a dense and well-adherent HA-chitosan
composite coating on AZ31 Mg alloy. They prepared HA-chitosan powder mixtures
containing up to 20 wt% of chitosan and commercial HA nanocrystalline powders
having a volumetric mean diameter of 15 nm and chitosan powders with a degree of
deacetylation of about 85% were used as the starting materials. To obtain powders
with an appropriate particle size, the as-received HA powders were heated at
1050
C for 2 h, and the chitosan powders were dry ball-milled for 12 h in a planetary
mill using ZrO
balls in a jar. The heat-treated HA powders were mechanically mixed
2
with the ball-milled chitosan po wders by dry ball milling and the HA-chitosan composite coating was deposited on AZ31 Mg alloy by aerosol deposition (AD), because
AD offered the advantage of room temperature deposition. Finally, various 5 mm-thick
Figure 3.17 Cross-section of the
DCPD-PCL coated Mg-Zn alloy.
Source: Wang et al. (2012) with
permission from Elsevier.
Surface design of biodegradable magnesium alloys for biomedical applications 113
HA-chitosan composite coatings were deposited on the Mg alloys. The composition
of the coatings was tailored by adjusting the HA and chitosan concentrations in the
powder mixtures. All the coatings exhibited high adhesion strength ranging from
24.6 to 27.7 MPa and better corrosion resistance than the bare Mg alloy. Moreover,
the biocompatibility of the coated alloy such as cell adhesion was improved appreciably as demonstrated by Figure 3.18.
Although Ca-P ceramics have favorable biocompatibility and osteoconductive
properties, they usually induce slow bone formation in vivo (Arinzeh, Tran, Mcalary,
& Daculsi, 2005). Compared to Ca-P ceramics, CaSiO
ceramics can promote prolif-
3
eration and differentiation of osteoblast-like cells and accelerate the formation of HA
in SBF, but, unfortunately, CaSiO
degrades rapidly in the physiological environment
3
(Siriphannon, Kameshima, Yasumori, Okada, & Hayashi, 2000; Ni, Chang, Chou,
& Zhai, 2007; Ni, Lin, Chang, & Chou, 2008). To improve the corrosion properties
and cell compatibility, Du et al. (2011) produced a microporous calcium silicate and
calcium phosphate (CaSiO
alloy by a chemical reaction. The layer was mainly composed of CaHPO
with a small amount of CaSiO
-CaHPO4$2H2O) composite coating on Mg-Zn-Mn-Ca
3
$2H2O
4
. In vitro cell experiments indicated that the surface
3
cytocompatibility of the coated Mg alloy was significantly improved as manifested
by more cell adhesion, growth, and proliferation.
Figure 3.18 SEM micrographs of the MC3T3-E1 cells attached to the samples: (a) Uncoated
AZ31 substrate, (b) HA coating, and HA-chitosan composite coatings with (c) 5 wt% and
(d) 20 wt% chitosan.
Source: Hahn et al. (2011) with permission from Elsevier.
114 Surface Modification of Magnesium and its Alloys for Biomedical Applications
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3.5 Summary and future trends
Biodegradability is a prominent advantage of Mg alloys in biomedical components
such as cardiovascular stents and bone fixation. Although this is a revolutionary
concept in biomaterials science contrary to traditional corrosion-resistant permanent biometals, there are several practical difficulties. The dynamic interface
between the Mg alloys and biological environment is quite c omplicated and
requires serious consideration. Rapid corrosion is a big issue especially in the
initial healing stage, and proper control is crucial. Therefore, it is imperative to
construct a temporary surface on Mg alloys to control the corrosion, improve
the biocompatibility, and preserve the mechanical performance during the healing
stage. Surface modification techniques including coating and ion implantation can
be conveniently used to alter selected features to address different clinical requirements. This chapter discusses some recent research activities on Ca-P based coatings, polymer-based coatings, MAO coatings, PVD coatings, and ion implantation
performed on Mg alloys. Most research activities have hitherto focused on in vitro
investigations, and many aspects of in vivo degradation have not been completely
understood.
Developing new biomedical Mg alloys is a main trend to substitute for traditional
alloys in some applications, and there are inevitably questions and problems. New
types of coatings or modified surface layers must be designed to meet practical requirements. In addition, hybri d surface treatment techniques will also be emphasized
because of the flexibility. With the aid of hybrid techniques, the temporary surface
can be endowed with more functions including the desirable drug controlled delivery
capability. More research is needed to better comprehend biodegradation in vitro and
in vivo in order to expedite clinical acceptance of the materials.
Acknowledgements
The work was supported by Hong Kong Research Grants Council (RGC) General Research
Funds (GRF) Nos. 112510 and 112210 and City University of Hong Kong Applied Research
Grants (ARG) Nos. 9667066 and 9667069.
References
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corrosion behaviour of magnesium alloys. Surface and Coatings Technology, 197,
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and Coatings Technology, 93, 158e167.
Arinzeh, T. L., Tran, T., Mcalary, J., & Daculsi, G. (2005). A comparative study of biphasic
calcium phosphate ceramics for human mesenchymal stem-cell-induced bone formation.
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