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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5846_Библиотеки_им_академика_М_И_Перельмана
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272 Surface Modification of Magnesium and its Alloys for Biomedical Applications
2. deposition of calcium phosphate coatings on magnesium and its alloys using a biomimetic
coating strategy, and
3. superhydrophobic coatings for corrosion control that mimic the topography and surface
chemistry of the lotus leaf.
10.1.1 General challenges in the development of coatings
for magnesium and its alloys
There are two key challenges in the development of coatings for magnesium alloys.
First, magnesium is a highly reactive metal that readily undergoes oxidation in
contact with air or aqueous solutions. This results in the formation of a magnesium
oxide/hydroxide layer at the surface. To further complicate matters, magnesium also
undergoes a reaction with atmospheric CO
surface. This Mg(OH)
/MgO/MgCO3surface layer can impede the adhesion and
2
uniformity of surface coatings deposited on these materials, resulting in poor coating
performance. A careful choice of pretreatment driven by the expected chemistry at the
magnesium alloy/coating bath interface is therefore essential for optimized coating
performance. Furthermore, magnesium reacts violently with most acids and readily
dissolves in aqueous media at pH values less than 11. This effectively eliminates
the use of many traditional coatings due to underlying dissolution of the substrate
during coating deposition, resulting in poor coating adhesion.
Second, the surface chemistry of these materials is not homogeneous. This is illus-
trated in Figure 10.1, which shows the surface microchemistry of three different magnesium aluminum alloys with 3%, 6%, and 9% aluminum, respectively. The b phase
(intermetallic, Mg
) precipitates at the grain boundaries during cooling of these
17Al12
alloys, resulting in a non-uniform surface chemistry. On the AZ31 alloy, this phase
is present as discrete bright spots in the image. As the amount of aluminum in the alloy
is increased, the b phase gradually becomes an interconnected network on the surface.
Therefore, the chemistry at the coating bath/alloy surface varies spatially across the
sample. Moreover, the surface chemistry of magnesium alloys varies dramatically
depending on the particular alloy to be coated. This means that coating strategies
are not n ecessarily applicable from alloy to alloy.
to produce magnesium carbonate at the
2
10.2 Modification of biodegradable magnesium alloy
implant surfaces
In recent years, magnesium alloys have been proposed as a potential biodegradable
metallic implant material for orthopedic and cardiovascul ar applications (Brar et al.,
2009; Staiger, Pietak, Huadmai, & Dias, 2006; Witte et al., 2008). They are an excel-
lent material for this purpose because they have mechanical properties that are similar
to bone (Staiger et al., 2006), have been shown to corrode in biological fluids
(Song et al., 2009; Xin et al., 2009; Yang & Zhang, 2009), and are inherently biocompatible (Brar et al., 2009; Witte et al., 2008; Staiger et al., 2006). Magnesium-based
metals readily corrode in aqueous solutions at pH values below 11; therefore, under
Biomimetic surface modifications of magnesium and magnesium alloys for biomedical applications 273
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Figure 10.1 Backscattered electron microscopy images of magnesium alloy AZ31 (top),
magnesium alloy AM60 (middle), and magn esium alloy AZ91 (bottom). The bright areas are
the aluminum-rich b phase of the alloy; the dark areas are the magnesium-rich a phase of
the alloy.
274 Surface Modification of Magnesium and its Alloys for Biomedical Applications
physiological conditions, these materials will readily degrade according to the
following reactions (Makar & Kruger, 1993):
0
(s) / Mg2þ(aq) þ 2e
Mg
2H2O (l) þ 2e/ H2(g)[ þ 2OH(aq)
Alloying with aluminum and zinc has been shown to improve the corrosion resistance of magnesium by increasing the stability of the oxide layer and the open circuit
potential (Anik & Celikten, 2007 ; Ghali, Dietzel, & Kainer, 2004a,b; Makar & Kruger,
1993; Mathieu, Rapin, Steinmetz, & Steinmetz, 2003; Song & Atrens, 1999; Song,
Atrens, & Dargusch, 1999). Alloying leads to the formation of both a magnesium-
rich phase (a phase) and an aluminum-rich intermetallic phase with a composition
of Mg
(b phase) in the alloyed material (Mathieu et al., 2003; Unigovski &
17Al12
Gutman, 1999). The corrosion resistance of magnesium alloys has been shown to
be related to the stability of the oxide layer that forms over the aluminum rich phase
of the material (Anik et al., 2006; Ani k & Celikten, 2007; Mathieu et al., 2003;
Song et al., 1999; Song & Atrens, 1999). Although alloying slows the corrosion
rate, hydrogen gas evolution and an increase in alkalinity are still by-products of the
corrosion reaction and pose significant health risks for the patient. Therefore, surface
modification of these materials to control the degradation rate in the early stages of
healing and improve their biocompatibility is crucial to the successful implementation
of magnesium alloy implants in medicine.
10.2.1 Surface modification to mimic the ECM
Cell adhesion and proliferation at the implant surface is a vital factor for successful
integration of a biomaterial within the body (Hersel, Dahmen, & Kessler, 2003). In
recent years, an improved understanding of the interactions of cells with the ECM
has led to the development o f coatings that mimic the structure and/or surface chemistry of the ECM. The ECM is an organized three-dimensional network of macromolecules with both micro and nanoscale surface features that provide cell support and cell
adhesion sites and regulate cell communication (Wang & Carrier, 2011). It is
composed primarily of polysaccharide chains of glycosaminoglycans and fibrous proteins such as collagen, elastin, and fibronectin. Several review articles detail biomimetic approaches for modulating cell adhesion on biomaterials (Avila et al., 2009;
Rahmany & Van Dyke, 2013 ). From these reviews, three general strategies for opti-
mizing cell adhesion, proliferation, and differentiation emerge: nanoscale patterning
of surfaces, the incorporation of bioactive molecules into surface coatings, and covalent bonding of bioadhesive ECM macromolecules/binding motifs to the biomaterials
surface.
Two research articles have highlighted the importance of nanoscale topography on
cell/magnesium surface interactions (Weng & Webster, 2012, 2013). In these articles,
the surface of magnesium ribbon was roughened by immersion in high concentration
sodium hydroxide solutions. The resulting surfaces were shown to have increased
Biomimetic surface modifications of magnesium and magnesium alloys for biomedical applications 275
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nanoscale surface roughness in comparison to the control sample. The increased
surface roughness was correlated to improved cell adhesion and cell function on the
modified surfaces. This is a promising area of future research; however, few studies
on the influence of magnesium surface topography on cell response have been
reported.
To date, two types of biomimetic surface modifications that mimic the ECM have
been the main focus of research efforts to improve both the corrosion resistance and
biocompatibility of magnesium alloys: biodegradable polymer coatings and covale nt
attachment of ECM proteins/binding motifs to magnesium alloy surfaces.
10.2.1.1 Biodegradable polymer coatings
Chitosan is a long-chain natural biopolymer obtained from chemical modificati on of
chitin, the main component of crustacean shells. This linear polysaccharide is made
up of randomly distributed b-linked
units, and it is similar in chemical composition to the glycosaminoglycans found in the
ECM. Chitosan has been shown to have good biocompatibility, which makes it an
attractive coating material for biomaterial applications. Several articles have been
published on the deposition and properties of chitosan coatings on magnesium alloys
(Bai et al., 2012; Gu et al., 2009; Hahn et al., 2011; Wu et al., 2010; Zhang, Dai,
Wei, & Wen, 2012 ). It has been observed that pure chitosan coatings produced via
dip-coating do not exhibit sufficient corrosion resistance for practical purposes
(Bai, et al., 2012; Gu et al., 2009). The topography and corrosion resistance of the coatings produced via dip-coating have been shown to be strongly influenced by the molecular weight of the chitosan polymer and the number of deposited layers. These
coatings vary from smooth and compact to porous at the microscale as a function of
coating procedure (Bai et al., 2012; Gu et al., 2009).
To overcome their limited corrosion resistance, micro-arc oxidation (MAO)
followed by deposition of a chitosan-based coating has been proposed. An MAO
magnesium phosphate layer was produced in one study, followed by pore sealing
with a dip-coated chitosan layer (Bai et al., 2012). The coatings produced by this combined method had better corrosion resistance than either MAO or chitosan coating
alone due to both the enhanced stability of the chitosan layer through cross-linking
reactions with surface phosphate ions and the barrier effect of the chitosan on the
microporous MAO layer (Bai et al., 2012). Several articles have reported the surface
modification of magnesium alloys through MAO followed by electrophoretic deposition of composite calcium phosphate/chitosan coatings (Wu et al., 2010; Zhang et al.,
2012). In these articles, it was demonstrated that the coating adhesion depended
strongly on the electrolyte composition and the voltage range employed.
Finally, aerosol deposition has also been proposed as a viable method for producing
hydroxyapatite (HA)-chitosan composite coatings on magnesium alloys (Hahn et al.,
2011). Aerosol deposition is a type of spray coating whereby dense, well ordered
ceramic coatings are formed by collision of solid particles of the coating material
with the substrate to be coated (Hahn et al., 2011). The deposition is done at room temperature, which allows this technique to be used to produce hybrid organic/inorganic
D-glucosamine and N-acetyl-D-glucosamine sub-
276 Surface Modification of Magnesium and its Alloys for Biomedical Applications
coatings. Aerosol deposi tion has been successfully used to produce HA-chitosan coatings on the AZ31 magnesium alloy. The topography of the coatings produced had a
rough surface, with a network structure that became denser and smoother as the concentration of chitosan in the coating feedstock increased. Improved corrosion
resistance and biocompatibility was observed for the composite coatings (Hahn
et al., 2011).
The development of biodegradable organic coatings on magnesium alloys using
several other polymers has also been reported. Alginate, a natural anionic polysaccharide that crosslinks to form a hydrogel upon exposure to calcium ions, has been
proposed as a potentially biocompatible coating for magnesium alloys (Sangeetha
et al., 2011). Although these coatings were shown to have good biocompatibility,
they were not particularly corrosion resistant due to their poor stability in vitro.
Similarly, dip-coated poly(lactide-co-glycolide) coatings were observed to signifi-
cantly enhance cell adhesion and proliferation but with limited improvement in
corrosion resistance due t o poor alloy-polymer adhesion (Ostrowskietal.,2013).
Spray-coated poly(lactic acid) and poly (DTH carbonate) films also yielded only
modest improvements for in vitro corrosion resistance (Gray-Munro, Seg ui n, &
Strong, 2009).
Other research has demonstrated that coating using polymers with functional
groups that can interact strongly with the magnesium alloy substrate, such as poly(vinyl acetate) (PVA), can result in improved adhesion and consequently improved corrosion resistance (Abdal-hay, Dewidar, & Lim, 2012). Furthermore, the influence of
morphology of the polymeric coating has also been shown to be an important factor
for optimum corrosion resistance and biocompatibility. The coating morphology is
strongly influenced by the solvent, polymer concentration, and method of deposition.
In one article, dip-coated PVA films were deposited on a magnesium alloy from
three different solvents (Abdal-hay et al., 2012). The coatings where the polymer
was dissolved in either tetrahydrofuran or dimethylformamide were a uniform, dense
layer, whereas those deposited from a dichloromethane/PVA solution were porous.
The porous coatings were observed to have the best corrosion resistance and improved
cell adhesion. Likew ise, spray-coated poly(lactic acid) coatings with a porous structure
were also found to have improved corrosion resistance with no cytotoxicity (Abdal-
hay, Barakat, & Lim, 2013b). Poly(lactic acid) coatings on magnesium alloys pro-
duced by dip-coating and electrospinning have also been compared (Abdal-hay
et al., 2013a). The dip-coated samples were shown to have a porous structure, while
electrospinning resulted in a fibrous structure similar to the ECM. Increased cell proliferation and decreased corrosion rates were observed on both coated samples. The
authors postulated that the porous structure prevents the build-up of hydrogen gas at
the substrate/coating interface, thus preventing the coating delamination that has
been observed for smooth dense polymer films.
The level of porosity in the biodegradable polymer fi lms has also been shown to be
an important factor for optimum in vivo performance. Porous poly(caprolactone) coatings with both high and low porosity have been reported (Wong et al., 2010). Both
coatings had improved corrosion resistance, excellent cell viability, and helped retain
the mechanical strength of the magnesium alloy implant during immersion testing.
Biomimetic surface modifications of magnesium and magnesium alloys for biomedical applications 277
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However, the low porosity coatings were observed to result in the most new bone formation, with no implant volume reduction for in vivo animal tests. In general, biodegradable polymer coatings have been show n to improve the corrosion resistance and
biocompatibility of magnesium and its alloys. However, future research efforts should
aim to improve the adhesion of these coatings and to better understand the role of
surface topography on corrosion resistance and biocompatibility.
10.2.1.2 Attachment of ECM proteins or peptides
Surface modification strategies that mimic the properties of the ECM through the
attachment of ECM proteins or peptide sequences present in ECM proteins have
been shown to dramatically improve the biocompatibility of implant surfaces (Hersel
et al., 2003; Rahmany & Van Dyke, 2013). Although this is a promising avenue of
research for magnesium alloys, it has not been extensively explored to date. One article
reported on the surface modification of the AZ31B magnesium alloy surface with a
peptide sequence designed to mimic the acidic bone protein, dentin sialophosphoprotein (Cui et al., 2013). The coating was deposited by dip-coating from a solution of the
peptide dissolved in tetrahydrofuran. This biomimetic peptide coating did have a
positive impact on the corrosion resistance of the substrate and was shown to facilitate
calcium phosphate precipitation in an organized crystal structure. Further research is
needed to optimize the corrosion resistance and to characterize the in vitro cell
behavior and in vivo tissue-regenerating ability of these coatin gs.
In the next section, preliminary results from my laboratory on two possible biomimetic surface modification strategies f or magnesium allo ys a re d isc uss ed. The first
is the covalent attachment of the arginine-glycine-aspartic acid peptide sequence
(RGD) to the magnesium alloy surface. Cells bind to surfaces through specific
ligands found in proteins of the ECMs. One of these ligands is RGD, commonly
found in adhesive proteins such as fibronect i n (Hersel et al., 2003; Rahmany &
Van Dyke, 2013). The bonding of peptide chains containing the RGD sequence to
biomaterials surfaces has been shown to significantly improve osteoblast attachment
to titanium and polymers (Chollet et al., 2009; Dettin et al., 2009). Furthermore, it
has been shown that factors such as the surface density and distribution of the
RGD peptide, the peptide sequence, the surface topography, and method of covalent
immobilization play a key role in mediating the observed cell response (Hersel et al.,
2003; Le Saux et al., 2011).
The second strategy is the covalent immobilization of collagen layers to the magnesium alloy surface. Bone tissue is a co mposite material consisting mainly of collagen
proteins and calcium phosphate. It is anticipated that chemisorbed collagen coatings
could provide a corrosion-resistant and biocompatible surface layer for improved
cell adhesion on magnesium alloys destined for orthopedic applications.
To achieve optimum stability of biomimetic coatings, covalent bonding to the
surface is essential. In this regard, organosilane coati ngs give a unique opportunity
to provide both the required corrosion resistance and the ability to biofunctionalize
the magnesium implant surface. These molecules can be synthesized with a wide
variety of functional groups at the end of the alkyl chain; an appropriate choice of
278 Surface Modification of Magnesium and its Alloys for Biomedical Applications
functional group will allow us to covalently attach biomolecules of interest. Two
papers have successfully demonstrated the attachment of protein molecules to the
surface of metallic magnesium using this strategy (Killian, Wagener, Schmuki, &
Virtanen, 2010; Wagener, Killian, Turhan, & Virtanen, 2013). The authors further
showed that their modified surfaces resulted in a significant decrease in the initial
corrosion rate of samples immersed in simulated body fluid (SBF).
Surface attachment of organosilane occurs when hydrolyzed organosilane
molecules undergo a condensation reaction with surface hydroxyl groups to form
Si-O-Metal bonds (Abel et al., 2006; Torry, Campbell, & Cunliffe, 2006). The relevant
reactions are given below, where R
0
is typically CH3or CH2CH3and X is a functional
group such as a thiol, alcohol, alkyl group, carboxylic acid, or amine:
X(CH
Si(OR0)3þ 3H2O / X(CH2)nSi(OH)3þ 3R0OH (Reaction 1:
2)n
Hydrolysis)
X(CH
Si(OH)3þ MOH / X(CH2)nSi(OH)2eOM þ H2O (Reaction 2:
2)n
Surface condensation)
Preliminary results from my research using organosilane coupling agents to covalently bond two different biomo lecules to the AZ31 magnesium alloy surface are
discussed in the next section.
Covalent attachment of the RGD peptide sequence through an
amino-functionalized organosilane
In this research, the RGD peptide motif has been covalently attached to the surface of
magnesium alloy AZ31 through an aminopropyltriethoxysilane (APTE S) layer using a
linker molecule, 3-maleimidopropionic acid n-hydroxysuccinimide ester (SMP), and
an RGD peptide sequence that contains a cysteine moiety. The SMP linker preferentially bonds to primary amines through its succinimide functional group and to thiols
through its maleimide functional group. This bifunctional crosslinker can thus covalently bond to the primary amine group of APTES at one end and to the cysteine
residue of the peptide at the other.
The Attenuated Total Reflection-Fourier Transform Infrared (ATR-FTIR) spectrum
for the APTES-coated magnesium alloy coupon is shown in Figure 10.2(a). This
spectrum shows the characteristic absorption bands expected for an APTES film.
The bands at 1136 cm
the presence of APTES on the surface (Kim & Seidler, 2009). The two peaks at
1475 cm
-1
and 1578 cm-1are due to the amine functional groups. The ATR-FTIR spectrum for APTES-coated AZ31 exposed to the linker molecule alone and then to the
linker molecule and the peptide are shown in Figure 10.2(b) and (c), respectively.
The infrared spectrum for pure SMP has a major infrared band at 1727 cm
pears as a weak shoulder at 1724 cm
due to eNH
peak (1578 cm-1) to lower wavenumbers (1566 cm-1) is indicative of inter-
2
action of the SMP molecules with the surface through the amine group. The ATR-FTIR
spectrum for the APTES-coated AZ31 after modification with both SMP and RGDC is
-1
(Si-O), 1044 cm-1(Si-O-Si), and 1408 cm-1(Si-CH2) confirm
-1
-1
(Figure 10.2(b)). Furthermore, a shift in the peak
. This ap-
Biomimetic surface modifications of magnesium and magnesium alloys for biomedical applications 279
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0.05
0.00
0.10
0.05
0.00
Absorbance
0.10
0.05
0.00
(c)
(b)
(a)
2000 1800 1600 1400 1200 1000
Wavenumbers (cm–1)
Figure 10.2 ATR-FTIR spectra of (a) APTES-coated AZ31, (b) SMP-modified APTES-coated
AZ31, and (c) RGD-modified APTES-coated AZ31.
shown in Figure 10.2(c). In this spectrum, the signals from the weaker bands due to
APTES and SMP in the 1400e750 cm
Si-O bands are still observed. Furthermore, two new strong vibrational bands at 1647
-1
cm
and 1586 cm-1are present on the surface. These are the characteristic amide bands
-1
region have been attenuated. Only the strong
associated with peptides. This indicates that the RGD peptide has been successfully
bonded to the surface. Sonication of the modified surface in water for 10 min did not
result in any change in the spectrum, indicating that the peptide is covalently bonded.
The influence of ATPES coating on the degradation rate of the magnesium AZ31
alloy in SBF was evaluated by comparing the surface chemistry and morphologies of
the uncoated and coated material after immersion in SBF for 24 and 72 h. The concentration of magnesium released into solution was also evaluated. The corresponding scanning electron microscopy (SEM) images are shown in Figure 10.3. It is clear that after
24 h immersion in SBF, there is a significant amount of pitting on the bare magnesium
surface (Figure 10.3(a)), while the APTES-coated surface shows no apparent evidence
of corrosion (Figure 10.3(c)). After 72 h, both samples (Figure 10.3(b) and (d))show
evidence of the deposition of a corrosion layer. However, the corrosion layer on the
bare magnesium surface appears to be thicker, as evidenced by the increased cracking
in the layer upon dehydration (Figure 10.3(b)) and the increased Ca/Mg ratio observed
by EDS (Figure 10.4). The increase in pH during corrosion results in precipitation of
insoluble calcium phosphate phases at the surface of magnesium alloys in SBF
(Gray-Munro & Strong, 2009). These calcium phosphate precipitates are calcium defi-
cient, as evidenced by the low Ca/P ratio observed (Figure 10.4).
Figure 10.4 also shows the overall amount of magnesium ions released into solution
for each sample. After 24 h, the amount of magnesium released is similar for coated
280 Surface Modification of Magnesium and its Alloys for Biomedical Applications
(a)
(c)
(b) (d)
Figure 10.3 Scanning electron microscopy images of samples after immersion in Hanks’
balanced salt solution: (a) Mg AZ31 after 24 h in SBF, (b) Mg AZ31 after 72 h in SBF, (c)
APTES-coated Mg AZ31 after 24 h in SBF, and (d) APTES-coated Mg AZ31 after 72 h in SBF.
and uncoated samples. However, after 72 h a noticeable decrease in the overall amount
of corrosion was observed for APTES-coated samples. This is likely due to defects in
the APTES layer that result in some localized corrosion of the alloy in the initial stages
of immersion. As these defects are filled in with insoluble calcium phosphates, the
corrosion rate slows. This indicates that the APTES films are stable in SBF, but that
film uniformity is an important factor in determining the overall corrosion resistance
of the mat erial.
Covalent immobilization of collagen
One of the main components of bone tissue is collagen proteins. The biocompatibility of
implant materials is strongly influenced by the nature of the protein layers that adsorb
upon implantation in the body because this layer governs cell response. To increase
biocompatibility, it is essential to control the nature of these adsorbed protein layers.
Proteins do not tend to adsorb to other proteins; therefore, one way to control the adsorption is to preadsorb a protein layer that will elicit a desired cell response. Because
Biomimetic surface modifications of magnesium and magnesium alloys for biomedical applications 281
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1.0
0.8
0.6
0.4
Ca/P ratio
0.2
0.0
0.016
0.012
0.008
0.004
Ca/Mg ratio
0.000
0.10
0.08
0.06
0.04
Mg (mg)
0.02
0.00
Mg 24 h
Mg 24 h
Mg 24 h
Mg 72 h
Mg 72 h
Mg 72 h
Sample
APS 24 h APS 72 h
APS 24 h
APS 24 h
APS 72 h
APS 72 h
Figure 10.4 Amount of magnesium released into solution, Ca/Mg EDS ratios, and Ca/P EDS
ratios for untreated magnesium alloy coupons and APTES-coated samples after immersion in
simulated body fluid for 24 and 72 h.
biodegradation followed by growth of new bone tissue is the target of magnesium alloy
implants, collagen coating may be one way to improve integration of these materials
into the body. In this project, collagen has been covalently attached to the surface of magnesium alloy AZ31 through a silane layer using a linker molecule (ascorbic acid), as
described by Killian et al. (2010), where albumin was covalently bonded to a magnesium
surface. The ascorbic acid contains two functional groups capable of bonding to primary
amines, thus enabling covalent bonding to the primary amine group of APTES at one end
and to primary amine containing amino acids on the protein molecule at the other.
Figure 10.5 shows the infrared spectra for APTES-coated AZ31 (the peak assign-
ments were discussed in Section 10.2.1.2.1) and APTES treated with ascorbic acid
(Figure 10.5(b)). Two important changes were observed. The first is that most of
the organosilane peaks are attenuated in comparison to the unmodified APTES layer,
confirming that the film is being covered. The second is the increase in relative peak
intensity of the peak at 1575 cm
amine groups interacting with HCO
-1
. This peak was previously attributed to protonated
. The increase in intensity is likely due to the
3
interaction of primary amine groups with ascorbic acid, which results in the formation
of a protonated amine species.
Figure 10.5(c) shows the infrared spectrum of the APTES-coated alloy modified
with ascorbic acid followed by dip coating in a collagen solution. The peaks due to
APTES are fully attenuated and two new peaks at z875 cm
-1
and 1400 cm-1are
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