Biofunctionalized anti-corrosive silane coatings for magnesium alloys

Acta Biomaterialia 9 (2013) 8671–8677

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Acta Biomaterialia

journal homepage: www.elsevier .com/locate /actabiomat

Biofunctionalized anti-corrosive silane coatings for magnesium alloys q

1742-7061/$ - see front matter � 2012 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.http://dx.doi.org/10.1016/j.actbio.2012.12.025

q Part of the Biodegradable Metals Conference 2012 Special Issue, edited byProfessor Frank Witte and Professor Diego Mantovani.

* Corresponding author. Tel.: +61 2 4298 1443; fax: +61 2 4221 3114.E-mail address: smoulton@uow.edu.au (S. Moulton).

Xiao Liu a, Zhilian Yue a, Tony Romeo a, Jan Weber b, Torsten Scheuermann c, Simon Moulton a,*,Gordon Wallace a

a ARC Centre of Excellence for Electromaterials Science, University of Wollongong, Wollongong, Australiab Boston Scientific Maastricht, Gaetano Martinolaan 50, 6229 GS Maastricht, Netherlandsc Boston Scientific Technology Center, Perchtinger Strasse 6, 81379 Munich, Germany

a r t i c l e i n f o

Article history:Available online 11 January 2013

Keywords:Magnesium alloyHeparinSilaneCorrosionBiodegradable metallic implants

a b s t r a c t

Biodegradable magnesium alloys are advantageous in various implant applications, as they reduce therisks associated with permanent metallic implants. However, a rapid corrosion rate is usually a hindrancein biomedical applications. Here we report a facile two step procedure to introduce multifunctional, anti-corrosive coatings on Mg alloys, such as AZ31. The first step involves treating the NaOH-activated Mgwith bistriethoxysilylethane to immobilize a layer of densely crosslinked silane coating with good corro-sion resistance; the second step is to impart amine functionality to the surface by treating the modifiedMg with 3-amino-propyltrimethoxysilane. We characterized the two-layer anticorrosive coating of Mgalloy AZ31 by Fourier transform infrared spectroscopy, static contact angle measurement and optical pro-filometry, potentiodynamic polarization and AC impedance measurements. Furthermore, heparin wascovalently conjugated onto the silane-treated AZ31 to render the coating haemocompatible, as demon-strated by reduced platelet adhesion on the heparinized surface. The method reported here is also appli-cable to the preparation of other types of biofunctional, anti-corrosive coatings and thus of significantinterest in biodegradable implant applications.

� 2012 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

1. Introduction

Metallic implants, such as stents, bone plates and artificialjoints, are widely used in the human body. However, there are seri-ous problems associated with permanent metallic implants,including restenosis, thrombosis, physical irritation, potentialinflammatory responses and an inability to adapt to the growthof and changes in the human body [1]. In some cases additionalsurgery is required to remove the implant after the tissue has suf-ficiently healed [2] and, as a consequence, researchers and clini-cians are looking towards biodegradable implants that, onceimplanted, only remain for an appropriate period to ‘‘fix’’ the prob-lem and then disappear [3]. Coupled with their strong mechanicalproperties and low cytotoxicity, magnesium alloys have attractedincreasing attention as candidate materials for biodegradablestents and bone plates [1,4–6]. However, the potential clinicalapplications of Mg alloys have been hindered by their poor corro-sion resistance. The rapid corrosion of Mg alloys results in the gen-eration of hydrogen bubbles and pH changes, which will damagesurrounding tissues [7]. More seriously, rapid corrosion can lead

to an early loss of mechanical stability of the Mg alloy implant be-fore the end of the healing process [4]. Therefore, an appropriatecorrosion rate becomes an essential requirement for the clinicalapplication of Mg alloys in biodegradable metallic implants.

Several techniques have been employed to improve the corro-sion resistance of Mg alloys, including the development of newMg alloys [8,9], surface modification via nitrogen ion implantation[10,11], anodizing [12], and the use of conversion coatings [13–15].Among these technologies, silane-based anti-corrosive coatings forMg alloys have been proven to be effective, economical andenvironmentally benign [16,17]. Silanes are a group of silicon-based organic–inorganic materials with the general formulaR0(CH2)nSi(OR)3, where R0 is an organofunctional group and R is ahydrolysable alkoxy group. When in contact with water silanesare hydrolysed to yield silanol groups (SiOH) that permit attach-ment to hydrated metal surfaces (metal-OH) via the formation ofSi–O–metal bonds [18]. The silanol groups undergo self-crosslinking via siloxane bonds (Si–O–Si), resulting in an organicprotective layer chemically bound to the metallic substrate [19,20].

Bistriethoxysilylethane (BTSE) and 3-amino-propyltrimethox-ysilane (c-APS) are a widely studied bis-silane and mono-silane,respectively. Such silanes can provide functional moieties that en-able further attachment of bioactive molecules to enhance theinterfacial interaction of metal implants with surrounding cellsand tissue [21]. However, it is reported that the amino group of

8672 X. Liu et al. / Acta Biomaterialia 9 (2013) 8671–8677

c-APS preferentially bonds to metal surfaces, which results in de-fects in the silane coatings, thus allowing water to penetrate tothe silane–metal interface [22,23]. Compared with functionalmono-silanes, non-functional ‘‘bis-silanes’’ provide better corro-sion protection due to the formation of densely crosslinkedthree-dimensional polysiloxane networks and stronger interfacialadhesion to various metal surfaces, including steel, Al alloys, Cu al-loys and Mg alloys [24]. To take advantage of both types of silanesa two step BTSE–c-APS coating treatment has been developed byVan Ooij’s group to provide good corrosion protection for alumin-ium and steel [25]. However, to our knowledge this two step silanecoating has not been applied to Mg alloys.

Another key requirement for biodegradable metallic implants,especially for cardiovascular stents, is blood compatibility. Adhe-sion of platelets can induce thrombus formation and, consequently,implant failure [26]. Heparin remains the most frequently usedanticoagulant reagent. Surface modification with heparin has beenintensively explored to increase the thrombo-resistance of bio-medical implants. Previous studies demonstrated that heparin-coated stents reduced stent thrombosis [27,28], and resulted infavourable event-free survival after 6 months [27].

In this study we have developed a two step BTSE–c-APS coatingstrategy to produce a biofunctionalized anti-corrosive coating onMg alloy AZ31. The silane layer was analysed using Fourier trans-form infrared (FTIR) spectroscopy and optical profilometry; theanti-corrosion properties of the coating were assessed by potentio-dynamic polarization and AC impedance measurements. We alsodemonstrated that bioactive heparin can be covalently attachedto the silane-modified Mg alloy surface. Using this technology wehave demonstrated both markedly improved corrosion resistanceand blood compatibility of the resulting Mg alloy.

2. Materials and methods

2.1. Materials

All reagents were used as received. BTSE, c-APS, N-hydroxy-succinimide (NHS), 1-ethyl-3-(3-dimethylaminopropyl) carbodi-imide (EDC), 2-morpholinoethanesulfonic acid (MES), heparin,phosphate-buffered saline (PBS), toluidine blue O (TBO) and glutar-aldehyde were from Sigma-Aldrich, Australia. AZ31Mg alloy sheetwith the nominal mass composition 96% Mg, 3% Al and 1% Zn waspurchased from Goodfellow Metals, UK. Simulated body fluid (SBF)containing 5.403 g NaCl, 0.504 g NaHCO3, 0.426 g Na2CO3, 0.225 gKCl, 0.230 g K2HPO4�3H2O, 0.311 g MgCl2�6H2O, 0.8 g NaOH,17.892 g HEPES, 0.293 g CaCl2 and 0.072 g Na2SO4 in 1000 mlMilli-Q water was freshly prepared.

2.2. Silanization of Mg alloy AZ31

The 2.0 mm thick AZ31Mg alloy sheets were cut into15 � 20 mm pieces and polished with progressively finer SiC pa-pers up to grit 2000. The samples were ultrasonically cleaned usingacetone, dried in air, and then immersed in a 3.0 M NaOH solutionfor 2 h to produce a uniform hydroxide layer on the substrates. TheNaOH-activated Mg substrates are referred to as Mg-OH.

BTSE or c-APS solution was prepared by mixing 5% silane, 90%ethanol and 5% Milli-Q water. The solutions were stirred at roomtemperature for 1 h to allow hydrolysis to proceed. The AZ31 sam-ples were then immersed in the hydrolysed BTSE solution at roomtemperature for 1 h, dried with hot air, and then cured at 120 �C for1 h. For the second step coating the BTSE-treated samples, denotedMg-B, were soaked in the c-APS solution at room temperature for30 min before being cured at 120 �C for 1 h. The resultant sampleswere denoted Mg-B-A.

2.3. Surface modification of Mg-B-A with heparin

EDC�HCl and NHS were added to a heparin solution(5.0 mg ml�1) in MES buffer to a final concentration of 2.0 mg ml�1.Mg-B-A samples were immersed in the above solution and shakenat room temperature for 4 h. The heparinized samples (Mg-B-A-heparin) were rinsed five times in both PBS and then water,respectively.

2.4. Physico-chemical characterisation of the modified AZ31 samples

The surface modified ZA31 samples were investigated using a Shi-madzu IRPrestige-21 FTIR spectrophotometer. Attenuated total reflec-tion (ATR) FTIR spectra were recorded at a resolution of 4.0 cm�1 overa scan range of 2000 to 700 cm�1. The surface morphology was ana-lysed using a Veeco optical profiler NT9000 (Veeco Instruments Inc.,USA). Static water contact angles were measured using the sessiledrop method (2.0 ll, Milli-Q water) with a Dataphysics OCA20 Goni-ometer (DataPhysics Instruments GmbH, Germany).

The amounts of surface accessible heparin were quantified usinga TBO assay [29,30]. The AZ31 samples were immersed in freshlyprepared TBO solution (0.04 wt.% in 0.01 M HCl/0.2 wt.% NaCl solu-tion, 2.0 ml per sample) and shaken gently at room temperature for4 h followed by rinsing five times with Milli-Q water. Then the AZ31samples were soaked in ethanol/NaOH (80/20 vol.%) solution for10 min, and the released TBO was quantified by measuring the opti-cal density of the solution at 530 nm. A standard curve was estab-lished using a series of standard solutions of heparin.

2.5. Anti-corrosion properties of the modified AZ31 sample

Potentiodynamic polarization curves were recorded at a scanrate of 5.0 mV s�1 in SBF using a CHI 660 system (CH InstrumentsInc., USA). Electrochemical impedance spectroscopy (EIS) measure-ments were carried out in SBF solution using a three electrode cellcomprising a 1.0 cm2 modified AZ31 working electrode, a platinummesh auxiliary electrode and an Ag/AgCl (3.0 M NaCl) referenceelectrode. The samples were placed in electrolyte and the open cir-cle potential monitored for 1 h. Impedance spectroscopy was con-ducted between 0.05 and 100 kHz at the measured open circuitpotential with an AC amplitude of 10 mV using a Gamry Potentio-static PCI 750 system (Gamry Instruments, USA). The long-termimmersion experiments were performed at 37 �C, the samples thenbeing removed and the electrochemical measurements performedat room temperature (24 �C).

2.6. Platelet adhesion analysis

Fresh whole rat blood, with EDTA as anticoagulant, was ob-tained from Australia Animal Resources Centre. The blood sampleswere centrifuged at 300g for 10 min at room temperature to isolateplatelet-rich plasma (PRP). After removal of the PRP, the remainingsamples were centrifuged at 2500g for 10 min at room tempera-ture to isolate platelet pool plasma (PPP). The number of plateletsin the PRP was diluted to 1 � 108 cells ml�1 by mixing PRP withPPP. Modified AZ31 samples were immersed in PRP for 1 h at37 �C. Thereafter, the samples were rinsed twice with PBS, fixedwith 2% glutaraldehyde for 2 h, dehydrated in a series of increasingconcentrations of ethanol to 100%, and then observed using a JEOL7500FA field emission scanning electron microscope.

2.7. Statistical analysis

All data are expressed as means ± SD, unless specified other-wise. An unpaired Student’s t-test was used for comparison, anda p value of less than 0.05 was considered to be statisticallysignificant.

Fig. 1. (A) Hydrolysis of BTSE; (B) hydrolysis of c-APS; (C) schematic illustration of the surface modification of AZ31Mg alloy.

Fig. 2. FTIR spectra of Mg, Mg-OH, Mg-B and Mg-B-A.

X. Liu et al. / Acta Biomaterialia 9 (2013) 8671–8677 8673

3. Results and discussion

3.1. Surface modification of the Mg alloy AZ31

The surface modification scheme for the AZ31Mg alloy is shownin Fig. 1. To facilitate silane coating, the AZ31Mg alloy was pre-treated with NaOH to generate surface hydroxide groups (Mg-OH). BTSE and c-APS were hydrolysed to yield silanol groups.Mg-OH was then sequentially treated with the hydrolysed BTSEand c-APS to introduce a two layer silane coating through oxanebond formation with the elimination of water. These processesare accompanied by in situ polycondensation of the hydrolysedBSTE and c-APS to form polysiloxane networks with surface aminefunctionality. In this study heparin was employed as an example ofa bioactive molecule, and conjugated to the aminized silane coat-ing of the AZ31Mg alloy to improve the blood compatibility.

The FTIR spectrum (Fig. 2) clearly shows peaks around 1045–1127 cm�1 that correspond to Si–O asymmetric stretching in–Si–O–Si– [19,25]. Subsequent coating with c-APS significantlyincreased the intensities of these peaks, with the appearance of anew peak around 1570 cm�1 that is assigned to protonated aminogroups [25]. Our FTIR results are consistent with those reported forBTSE- and/or c-APS-treated Al, Fe and Mg samples [17,19,25,31].

3.2. Physico-chemical properties of the surface modified AZ31Mgalloys

The influence of modification on the surface morphology of theMg alloys was investigated (Fig. 3). The bare AZ31 sample showed

a uniformly patterned morphology associated with the polishingprocedure, with a surface roughness (Ra) of 148.7 ± 16.5 nm (Ta-ble 1). Mg-OH still retained the polishing-induced regular pattern,but with an increased Ra of 203.1 ± 6.5 nm. A much rougher surfacemorphology (Ra � 910.3 ± 2.2) was noted for Mg-B, attributable tothe BTSE coating. The subsequent modification with c-APS mark-edly reduced the Ra of Mg-B-A to 761.7 ± 51.2 nm, is presumably

Fig. 3. Surface morphologies of AZ31 samples following different treatments: fresh polished (A) Mg, (B) Mg-OH, (C) Mg-B, (D) Mg-B-A and (E) Mg-B-A-heparin.

Fig. 4. Static water contact angles of Mg, Mg-OH, Mg-B, Mg-B-A and Mg-B-A-heparin samples.

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due to the presence of grafted c-APS as an overlayer on the modi-fied AZ31 surface. Further conjugation of heparin to Mg-B-A re-sulted in the roughest surface morphology amongst the testedsamples, with a Ra of 2320 ± 57.7 nm. During the two step silanecoating and heparin modification the polishing-induced regularpattern that is evident on bare AZ31 becomes less distinctive.

The influence of modification is also reflected in the changes insurface wettability of the modified AZ31 alloys, as characterized bythe static contact angle measurements (Fig. 4). Bare AZ31 exhibitssurface hydrophobicity, with a contact angle of 75.2 ± 2.0�. Thepretreatment with NaOH increased the surface hydrophilicity ofMg-OH, with a water contact angle of 14.4 ± 1.9�. As expected,incorporation of the BTSE layer re-established the surface hydro-phobicity, with the static contact angle increasing to 103.2 ± 1.8�.Compared with Mg-B, Mg-B-A is less hydrophobic, with a lowercontact angle of 78.2 ± 1.5�, ascribed to the presence of surfaceamine moieties. Further modification with heparin rendered thesurface hydrophilic [32], as characterized by a much lower watercontact angle of 21.4 ± 2.2�.

3.3. Corrosion resistance properties of the modified AZ31 Mg alloys

Functional silanes, such as c-APS, have often been used to mod-ify stainless steel [33] and Ti [30,34] to provide surface functional-ity for post-chemical anchorage of biomolecules (e.g. collagen,heparin and fibronectin), with a view to improving the biocompat-ibility of these metallic implants. However, functional silanes, suchas c-APS, have failed to provide effective corrosion resistance tometals and their alloys [24]. As reported by others [24,25], thereare two regions within a silane coating which can provide corro-

Table 1Contact angle, surface roughness (Ra), corrosion current (Icorr) and corrosion resistance (Rp

AZ31 sample Abbreviations Ra (nm)

As-polished Mg 148.7 ± 16.5NaOH passivated Mg-OH 203.1 ± 6.5BTSE-coated Mg-B 910.3 ± 2.2BTSE + APS-coated Mg-B-A 761.7 ± 51.2BTSE + APS + heparin-coated Mg-B-A-heparin 2320 ± 57.7

The sample size (n) for the roughness and contact angle measurements was 5 and 3 for

sion resistance: the crosslinked outer silane layer enriched withSi–O–Si bonds and the interfacial layer dominated by Si–O–metalbonds. The more hydrophobic the outer layer, the lower the waterpenetration rate [25]. As a bis-silane such as BTSE is more hydro-phobic than c-APS, and the interfacial regions between a bis-silaneand a metal contains a higher density of Si–O–metal bonds thanthat of a mono-silane, BTSE provides better corrosion protectionthan c-APS in both the outer layer and the interfacial layer. Silaneshave been widely used as coupling agents in clinical applicationsfor more than 50 years, particularly in dentistry as adhesion pro-moters [35]. They have proven to be safe in vivo [35,36]. However,the applications of silane coatings in dentistry are limited by bond

) values of the Mg alloy samples after different treatments.

Contact angle (�) Icorr (lA cm�2) Rp (ohm cm�2)

75.2 ± 2.0 8.32 ± 0.63 2650 ± 53814.4 ± 1.9 5.10 ± 1.42 3178 ± 787

103.2 ± 1.8 2.69 ± 0.31 7788 ± 257278.2 ± 1.5 0.90 ± 0.24 13635 ± 274521.4 ± 2.2 1.87 ± 0.23 9515 ± 497

Icorr and Rp.

Fig. 6. Impedance spectra of Mg, Mg-OH, Mg-B, Mg-B-A and Mg-B-A-heparinsamples.

Fig. 7. Impedance spectra of Mg and Mg-B-A-heparin samples in SBF solution after1, 6, 12, 18, 24, 30, 36 and 42 h.

X. Liu et al. / Acta Biomaterialia 9 (2013) 8671–8677 8675

degradation [35], associated with hydrolytic cleavage of the silox-ane bonds. In degradable implant applications the purpose of thesilane coating is to slow down the initial corrosion rate, rather thantotally stop the corrosion process. Thus bond degradation is ex-pected when water penetrates to the interface between the silanecoating and the Mg substrate, resulting in detachment of the silanecoating. Previous studies have shown that stents will be encapsu-lated by new intimal tissue after implantation [37], therefore, aslong as detachment occurs after new intima formation, silane coat-ing fragments may be retained and localized rather than being lostto the bloodstream. Thus the risks associated with silane coatingdetachment can be controlled by the development of novel silanecoupling agents with increased bond strength and hydrolyticstability. The two step coating procedure reported here aims tocombine the advantages of both BTSE and c-APS to produce anti-corrosive coatings with surface functionality. The chemicallybound BTSE layer introduced in the first coating step serves as ahydrophobic barrier, improving the hydrolytic stability of the si-lane coating and protecting the underlying Mg from rapid corro-sion. This is clearly demonstrated in the corrosion resistancestudy (Figs. 5 and 6) of the modified Mg alloys.

Polarization curves for the modified AZ31Mg alloys are shownin Fig. 5 and the corrosion current (Icorr) values are summarizedin Table 1. Compared with bare Mg, the Icorr value decreased by�68% for Mg-B, and �89% for Mg-B-A. As Icorr is directly propor-tional to the corrosion rate [38], a distinct improvement in corro-sion resistance is indicated. The corrosion of magnesium involvesreaction of H2O with Mg to produce Mg(OH)2 and H2. The overallreaction includes anodic (Mg ? Mg2+ + 2e�) and cathodic(2H2O + 2e�? H2 + 2OH�) reactions [39]. After NaOH passivationthe AZ31 surface was covered with a layer of Mg(OH)2 film. Metalcation transport dominates the anodic kinetics as an anodic Mgdissolution reaction occurs on the AZ31 surface underlying theMg(OH)2 film. The Mg(OH)2 film cannot stop electrolyte penetra-tion as it is porous and unstable in an aqueous environment. Thecharge transfer associated with the cathodic reaction can occurboth beneath and on top of the film, resulting in a higher cathodiccurrent density compared with the silane-coated samples. AfterBTSE coating the exchange current density due to the cathodicreaction is reduced because the hydrophobic silane film acts as aphysical barrier retarding water penetration and electrontransport. The anodic dissolution reactions rate is also slowed asthe silane coating blocks mass transport of Mg2+ [24]. At the sametime, the formation of Si–O–Mg bonds at the interface also blocks

Fig. 5. Polarization curves of Mg, Mg-OH, Mg-B, Mg-B-A and Mg-B-A-heparinsamples.

Fig. 8. Quantitative characterization of the surface accessible heparins on Mg-B-Aand Mg-B-A-heparin samples.

some anodic reactions [24,31]. The second coating layer of c-APSreduces electrolyte penetration and electron transport, which fur-ther reduces the corrosion current by around one order of magni-

Fig. 9. Representative SEM micrographs of platelets on (A) Mg-B-A and (B) Mg-B-A-heparin. The scale bars represent 2 lm.

8676 X. Liu et al. / Acta Biomaterialia 9 (2013) 8671–8677

tude compared with uncoated AZ31 samples. The shift in corrosionpotential towards the cathodic direction also indicates that thehydrophobic silane film acts as a physical barrier to retard electro-lyte penetration [31]. Further modification of Mg-B-A with heparinresulted in an increase in Icorr from 0.9 to 1.87 lA, which may bedue to the outer layer silane film being saturated with electrolyteduring the covalent conjugation process and thus partly losing itsfunction as a physical barrier to retard electrolyte penetration[24]. However, even in this case, the Icorr of Mg-B-A-heparin is stilllower than that of bare AZ31.

The anti-corrosive properties of the modified AZ31 sampleswere also assessed by EIS (Fig. 6). Over the whole frequency rangeexamined applying a BTSE or BTSE–c-APS coating resulted in a pro-gressive increase in impedance. Compared with bare AZ31, theimpedance of Mg-B-A increased by one order of magnitude as a re-sult of the non-conductive silane coating, which indicates effectiveprotection of the metal surface [24]. Similarly to the Icorr result, theimpedance of the Mg-B-A-heparin sample was reduced to a similarlevel to that of Mg-B, but was still four times higher than that ofthe untreated AZ31 sample. The EIS spectra of the Mg and Mg-A-B-heparin samples were also recorded in SBF at 37 �C every 6 hover a 42 h period (Fig. 7). Two capacitive semicircles are clearlyobserved over the first hour for the uncoated AZ31 and silane-coated samples. For the uncoated AZ31 samples the middle fre-quency (mf) semicircles disappeared and a low frequency (lf)inductive loop appeared after 6 h immersion, indicating severepitting corrosion [40]. For the silane-coated samples (Mg-B-A-heparin) both semicircles were clearly evident up to 24 himmersion, after which time both semicircles appear to merge intoone large semicircle with an associated increase in impedance,suggesting improved corrosion protection [17,41].

3.4. Platelet adhesion to the modified AZ31 Mg alloys

Compared with the single functional silane coating approach re-ported previously [16,17,42], this work utilized BTSE as the firstcoating layer to provide a hydrophobic barrier while minimizingthe interference by functional groups (e.g. amino groups [23]),which is important for the improved corrosion resistance of mod-ified AZ31. In addition, the unique two step coating offers the po-tential for further modification to improve the biocompatibility ofMg alloy implants. In this work heparin was covalently conjugatedto the modified Mg surface via the amino groups provided by thegrafted c-APS. The density of surface accessible heparin, as deter-mined by the TBO assay, was �12 lg cm�2 (Fig. 8), which is signif-icantly higher than that of the blank control. This value iscomparable with those of heparin-modified surfaces, which wereshown to effectively improve blood compatibility [30,43]. To assesswhether the heparin retained bioactivity after covalent conjuga-tion, platelet adhesion assays were performed across different

samples. In contrast to the silane-modified AZ31 samples, signifi-cantly lower platelet adhesion was observed on the Mg-B-A-hepa-rin surface (Fig. 9).

4. Conclusions

Through a two step coating process we have developed biofunc-tionalized anti-corrosive silane coatings for biodegradable Mgalloys. Compared with bare AZ31Mg alloy, Mg-B-A-heparin exhib-its both an improved corrosion resistance and reduced plateletadhesion. The development of a surface modification strategy thatcan simultaneously control the Mg alloy corrosion resistance andinhibit platelet adhesion will inevitably have major significancefor improving the blood compatibility of biodegradable metallicimplants. In addition, this work opens up new avenues and apotential platform to functionalize Mg alloys based on chemicalanchorage and in situ condensation of two types of silanes, a pro-cess which then provides functional groups for the further immo-bilization of essential biological components to facilitate theongoing development of biodegradable metallic implants.

Acknowledgements

The authors would like to acknowledge Boston Scientific andthe ACES for funding through ARC linkage Grant LP0990621. Thisresearch used equipment funded by the Australian ResearchCouncil and located at the UOW Electron Microscopy Centre andAustralian National Fabrication Facility – Materials Node. Specialthanks are due to Mr. Darren Attard of the UOW ElectronMicroscopy Centre and Dr. Stephen Beirne of the IntelligentPolymer Research Institute for assistance in sample preparation.

Appendix A. Figures with essential color discrimination

In this article, Fig. 3 is difficult to interpret in black and white.The full color images can be found in the on-line version, athttp://dx.doi.org/10.1016/j.actbio.2012.12.025.

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