磷酸盐6

th rt d sity d fo e 1 Abstract rials have limitations that can be clinically detrimental. For instance, the elastic modulus of currently used metals is stress shielding effects [3]. This is typified by increased bone Certain coatings on metallic biomaterial surfaces have been shown to improve corrosion resistance and improve the bioactivity of the surface through osteoconduction, i.e. bone ingrowth [8–10]. Several different coating approaches have been investigated to change the biological properties of metals and improve osteoconduction. These methods * Corresponding authors. Tel.: +1 5143987203x089737 (F. Tamimi); tel.: +1 514 3983908 (J.E. Barralet). E-mail addresses: faleh.taminimarino@mail.mcgill.ca (F. Tamimi), jake.barralet@mcgill.ca (J.E. Barralet). Available online at www.sciencedirect.com Acta Biomaterialia 5 (2009) 2338–2347 Metals have been the biomaterials of choice for biomed- ical load-bearing applications due to their combination of high mechanical strength and fracture toughness [1]. Hence, they are widely used as screws, plates, pins and implants (both orthopaedic and dental) where bone stabil- ization and/or augmentation are required. The current types of metals used are cobalt chromium (molybdenum) alloys, stainless steel 316L, pure titanium and titanium alloys [2]. Despite sufficient mechanical strength, metallic biomate- resorption around an implant as a disproportionate amount of the load is taken by the metal rather than the surrounding bone. Since the peri-implant bone undergoes resorption, the implant may loosen and eventually fail [3]. Additionally, metallic surfaces have a limited capacity for integration with bone, which is determined by features such as surface topography and chemistry, and can release toxic ions through corrosion or mechanical wear [4,5]. This can stimulate an inflammatory response and subsequently loss of bone, which can lead to implant loosening [6,7]. Osteoconductive coatings may improve the clinical performance of implanted metallic biomaterials. Several low-temperature coating methods have been reported where a supersaturated solution is used to deposit typically apatitic materials. However, due to the very low solubility of apatite, the concentration of calcium and phosphate ions attainable in a supersaturated solution is relatively low (�1– 2 mM), thus coating formation is slow, with several solution changes required to form a uniform and clinically relevant coating. In order to avoid this problem, we present a novel method where substrates were initially sputter coated with pure magnesium metal and then immersed in differing phosphate solutions. In this method, upon immersion the implant itself becomes the source of cations and only the anions to be incorporated into the coating are present in solution. These ions react rapidly, forming a continuous coating and avoid- ing problems of premature non-localized precipitation. The different coatings resulting from varying the phosphate solutions were then characterized in terms of morphology and composition by microscopy and chemical analyses. Upon immersion of the sputter-coated metals into ammonium phosphate solution, it was found that a uniform struvite (MgNH4PO4�6H2O) coating was formed. Upon subse- quent immersion into a calcium phosphate solution, stable coatings were formed. The coated surfaces also enhanced both osteoblastic cellular adhesion and cell viability compared to bare titanium. The concept of sputter-coating a reactive metal to form an adherent inor- ganic metal coating appears promising in the field of developing rapid-forming low-temperature bioceramic coatings. � 2009 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. Keywords: Titanium coating; Magnesium sputtering; Struvite; Surface characterization; Osteoblast 1. Introduction much higher than that of natural bone, and this can create Magnesium-sputtered titanium for Suzette Ibasco, Faleh Tamimi *, Robe Srikar Vengallatore, Edwar Faculty of Dentistry, McGill Univer Received 5 January 2009; received in revise Available onlin 1742-7061/$ - see front matter � 2009 Acta Materialia Inc. Published by Else doi:10.1016/j.actbio.2009.03.006 e formation of bioactive coatings Meszaros, Damien Le Nihouannen, Harvey, Jake E. Barralet * , Montreal, QC, Canada H3A 2B2 rm 3 March 2009; accepted 5 March 2009 4 March 2009 www.elsevier.com/locate/actabiomat vier Ltd. All rights reserved. ate include plasma spraying of hydroxyapatite, alkali treatment of titanium surfaces to induce mineralization and direct pre- cipitation of apatites in simulated body fluid [11,12]. Plasma spraying is themost commonly usedmethod for coatingmet- als [13]. However, this is a high-temperature process which limits the application to thermally stable coatings and sub- strates. Also, plasma-spraying does not allow for the coating of geometrically complex andporous surfaces [13].Although aqueous deposition methods offer a solution to the limita- tions of plasma spraying, being a low-temperature coating process and able to coat any exposed surface, the technique is time consuming and yet to be commercially realized [14,15]. Alkali treatment of titanium [16,17] has performed well in vivo; however, the scope for varying surface and sub- strate chemistry is limited. Low-temperature coating meth- ods are very attractive since they allow the incorporation of thermally unstable biologically active compounds such as growth factors, e.g. bone morphogenic proteins [18], adhesionmolecules [19] and antibiotics [20] thatmay be used to improve the clinical performance of the metallic implant. Traditionally, aqueous solution deposition of hydroxyap- atite coatings is slow (in the order of days) [21]. This is due to the very low solubility of apatite such that the concentration of calciumand phosphate ions attainable in a supersaturated solution is relatively low (�1–2 mM). Therefore the solu- tion’s capacity to act as a simultaneous source of anions and cations is limited, slowing the deposition rate and requir- ing continuous renewing of the solutions [21]. We sought to avoid this problem by providing a source of cations on the metallic surfacewithwhich to react and forma calciumphos- phate layer. To achieve this, the metallic substrate was first sputter-coated with pure magnesium metal, then immersed in differing phosphate solutions to effectively ‘‘pre-corrode” the magnesium before substituting the magnesium phos- phate with calcium phosphate. Continuous coatings were formed rapidly therefore avoiding problems of premature non-localized precipitation. Interestingly, magnesium phosphates may form patho- logically in the human body and are found mainly in the form of struvite crystals (MgNH4PO4�6H2O) in kidney stones through the precipitation of trivalent phosphate in combination with ammonium ions [22–24]. Recent studies have exploited the biological importance of this mineral by suggesting a biodegradable and osteoconductive stru- vite-based cement for bone regeneration procedures [25]. However, despite promising reports, the magnesium phos- phates remain relatively unstudied as bioceramics. This study investigated the use of sputter coatings of pure mag- nesium metal on titanium substrates as a solid phase reac- tant for the production of inorganic bioactive coatings. 2. Materials and methods 2.1. Magnesium sputtering of metallic substrates S. Ibasco et al. / Acta Biom Titanium sheet samples (6Al–4V alloy; 25.0 mm � 25.0 mm � 0.5 mm; McMaster-Carr Company, Los Ange- les, CA, USA) were initially ultrasonically rinsed for 15 min in a 50/50 wt.% ethanol/acetone solution. Without any further pre-treatment, the titanium alloy sheets were sputter coatedwith theDentonExplorer�-14 sputter coating system with a 3 lm thick magnesium layer using 99.5 wt.% magnesium metallic targets (Goodfellow Cambridge Lim- ited, England) of 20 cm diameter and a power of 150 W in high purity argon. 2.2. Coating of metals In order to determine which solutions would be suitable for the formation of intact coatings, experiments were first performed by incubating the magnesium-sputtered titanium sheets in simulated body fluid solution (SBF) [26] and in an SBFwhere potassium phosphate was substituted for sodium phosphate (Table 1). Next, SBF components were examined individually to determine which, if any, were capable of forming a coating. The four that formed a coating, as deter- mined by scanning electron microscopy (JEOL JSM-840A scanning electron microscope with an energy-dispersive X- ray (EDX) detector, both operating at 10–15 keV) were mixed in all permutations to examine the combinational effect on coatingmorphology. Only the three- and four-com- ponent combinations that produced any coating are pre- sented in Table 1. The magnesium-sputtered samples were immersed in different precipitation solutions with a pH of 7.4 at 37 �Cfor 24 h (Table 1). In the case of struvite coatings, magnesium-coated samples were also immersed in ammo- nium diphosphate solution for 30 s and 2, 15 and 120 min, the reactions were arrested by immersing the samples in eth- anol 100% and then vacuum dried before further character- ization. Volumes (Vs; ml) for the immersion liquid were calculated based on the formula: V s ¼ Sa=10 where Sa (cm 2) was the surface area of the sample as de- scribed previously [27]. All coatings that were rapidly soluble in both phosphate- buffered saline (PBS) and foetal bovine serum (FBS) were stabilized by immersion in a dilute calcium phosphate solu- tion for 48 h (Table 1). The non-sputtered reverse side of the titanium sheets was used as a negative control. 2.3. Characterization of the coatings All metal coatings were sputter coated with Au/Pd before being examined under scanning electron microscopy (SEM) coupled with energy dispersion spectroscopy at a potential of 10 kV and a working distance of 12.0 mm to determine their morphology and elemental composition. X-ray diffraction (XRD) and atomic force microscopy (AFM) analysis of the coatings were performed to evaluate their crystallographic nature and topology. A vertical-goni- ometer X-ray diffractometer (Philips model PW1710, rialia 5 (2009) 2338–2347 2339 Bedrijven b. v. S&I, The Netherlands), equipped with a Cu Ka radiation source, was used for the powder diffrac- m d ate tion pattern collection. Data were collected across a 2h range of 10–80�, with a step size of 0.02� and a normalized count time of 1 s per step. The phase composition was examined by means of the International Centre for Diffrac- tion Data (ICDD) reference patterns. An AFM (MFP-3D- BIO/Olympus IX71; Asylum Research, Santa Barbara, CA) in non-contact mode was used to make high-definition topographical images of the dry surfaces of the samples at room temperature. In order to investigate coating thickness and homogene- ity, coated metals were embedded in a resin mounting med- ium (Technovit 2000 LC, Heraeus Kulzer GmbH, Wehrheim Germany). The samples were then cut and ground through the metal coating with a silicon carbide grinding sequence (600, 800, 1200 and 2400 grits), followed by polishing with a fine alumina slurry on a medium nap cloth an examination under SEM. Durability of the coating was assessed using a peel test [27]. To test the stability of the coatings under physiologi- cal conditions, coated samples were immersed in PBS and FBS for 24 h at 37 �C and samples were examined by Table 1 Composition and concentrations (mM) of the reacting solutions. Reactants Sample ID Modified SBF A B C KCl 0.225 0.225 – Na2HPO4�7H2O – 0.231 0.231 MgCl2�6H2O 0.311 – 0.311 K2HPO4�3H2O 0.231 0.231 0.231 NaCl – – – KH2PO4 – – – NH4H2PO4 – – – CaCl2 – – – NaHCO3 – – – Na2SO4 – – – * SBF, Kokubo’s SBF [26]; PBS, phosphate-buffer saline; ADP, ammoniu 2340 S. Ibasco et al. / Acta Biom SEM and AFM. 2.4. Osteoblast cell culture The murine pre-osteoblastic cell line MC3T3-E1 was obtained from the American Type Culture Collection (ATCC, Rockville, USA). MC3T3-E1 cells were cultured in 25 cm2 tissue culture flasks in MEM culture medium with 10% FBS, 1% L-glutamine, 1% penicillin/streptomycin and 1% sodium pyruvate at 37 �C in a humidified atmo- sphere containing 5% CO2 in air. Cells were sub-cultured once a week using trypsin/EDTA and maintained at 37 �C in a humidified atmosphere of 5% CO2 in air. The medium was renewed every 2 days. 2.5. Cell adhesion The MC3T3-E1 cell adhesion was quantified on square samples (1 cm2) of bare titanium (T) and calcium phos- phate-coated titanium (CT). Cells were seeded onto the surface of T and CT in 24-mutiwell plates at a final density of 2.0 � 106 cells cm�2. After 2 h, non-adherent cells were removed from the material surfaces by vortexing the sam- ples in fresh culture medium for one minute. The attached cells were determined by cell counting using a hemocytom- eter following cell detachment by treatment with a trypsin/ EDTA solution (0.05%). Experiments were performed in triplicate. 2.6. Cell viability The viability of pre-osteoblastic cells was determined by the Live/Dead� Viability/Cytotoxicity Kit (Molecular Probes, Eugene, OR, USA). This kit provides a two-color fluorescence cell viability assay which is based on the simul- taneous determination of live and dead cells by detection of the intracellular esterase activity by calcein AM and of plasma membrane integrity by ethidium homodimer (EthD-1), respectively. Cells were washed three times with sterile PBS and incubated for 30 min at room temperature SBF* PBS ADP CaP D E F G H 0.225 0.225 2.68 – – 0.231 – 8.10 – 1.34 0.311 0.311 – – – 0.231 0.231 – – – – 8.035 136.89 – – – – 1.47 – – – – – 500 – – 0.292 – – 2.00 – 0.355 – – – – 0.072 – – – ihydrogen phosphate (NH4H2PO4); CaP, calcium phosphate. rialia 5 (2009) 2338–2347 in 20 ll of PBS solution containing 2 lM calcein-AM and 4 lM EthD-1. The samples were then examined under a fluorescence microscope (Eclipse TE 2000-U, Nikon Instruments Inc., Melville, NY, USA) (�4). The percent- age of viable cells was determined by scanning five ran- domly chosen fields from each slide, with at least 100 cells being analyzed per field. 3. Results SEM examination revealed that aqueous solutions con- taining KCl, Na2HPO4�7H2O, MgCl2�6H2O, K2HPO4, or NH4H2PO4 as a single component reacted with the magne- sium layer to form stable phosphate precipitates, while no precipitates were formed on the non-sputtered side of the titanium sheets. Although precipitates did form when immersed in these single component solutions and two component combinations of the same salts, the coating was patchy at the microscopic level (not shown). Three-component combinations were slightly better than the ones prepared with two components (Fig. 1, samples A–C), and a uni- formly cracked magnesium phosphate coating was appar- ent when combining all four active components (Fig. 1, sample D). In Table 1 we summarize the composition of the solutions containing three and four components that formed any insoluble precipitates on the magnesium-sput- tered titanium. It was also determined by EDX analysis that agglomer- ated precipitates of magnesium phosphate and magnesium chloride could be achieved in immersing solutions contain- ing phosphate and chloride ions. Immersion in both non- modified SBF solution (Fig. 1, sample E) and PBS solution showed that although a coating of magnesium phosphate was formed, it was irregular in morphology and cracked in places (Fig. 1, sample F). Following immersion in ammonium diphosphate solution (ADP), struvite crystals rapidly formed over the surface within a few seconds, to be fully coated in just 2 min (Fig. 2). After 24 h of immer- tals and formed a crack-free surface (Fig. 3B). EDX line scan analysis of the cross-section indicated the thickness of the calcium phosphate layer to be �10 lm (Fig. 4). Lin- ear EDX confirmed the XRD analysis by revealing that the coating composition is of calcium phosphate (Fig. 4). EDX analyses revealed that the coatings formed on samples A–F were thin and/or cracked, as strong peaks for the metal substrate were found even at low beam cur- rent (data not shown). However, EDX analysis of the stru- vite coating (Fig. 3C) confirmed the presence of a thick uniform magnesium phosphate coating. Furthermore, the composition of this coating changed into calcium phos- phate when immersed in CaP solution (Fig. 2D). XRD analysis confirmed the presence of struvite in the samples incubated in ADP solution. Residual unreacted magnesium metal was also detected (Fig. 3E, sample G). XRD patterns of struvite coatings after the secondary coat- ing inCaP solution revealed peaks characteristic of hydroxy- apatite, and no magnesium metal was detected (Fig. 3F). S. Ibasco et al. / Acta Biomaterialia 5 (2009) 2338–2347 2341 sion in ADP, the dense and uniform coating of struvite appeared to be stable without any apparent cracks (Fig. 3A, C and E). The struvite crystals appeared to radi- ate from a central nucleus forming circular polygonal microscale domains across the surface (Fig. 3A). Linear EDX of the coating cross-section, as well as the XRD, revealed that a small amount of magnesium remained unre- acted below the struvite coating adhered to the titanium surface (Fig. 3E). Interestingly, linear EDX analysis of the coating cross-section revealed small amounts of magne- sium and oxygen between the external struvite layer (Fig. 4) and the inner unreacted Mg; this could correspond to an intermediate layer of magnesium hydroxide. When com- pared to magnesium-sputtered metals immersed in CaP solution, no coatings were achieved, only bare metal. Post-treatment of the struvite coating in a CaP solution showed a uniform array of calcium phosphate platelet crys- Fig. 1. (A–C) SEM images of magnesium coated metals immersed in three-co and (F) PBS. Precipitated struvite formed block-shaped microcrystals with heights up to �2 lm that were imaged using AFM (Fig. 5A). Upon immersion in a calcium phosphate solu- tion the struvite microcrystals were replaced with thin (100–200 nm) plate-like crystals up to 2.5 lm in height and between 2 and 4 lm in length (Fig. 5B). The struvite and CaP coatings were durable enough to withstand the peel test as determined by SEM examination before and after the test (not shown). Attachment of pre- osteoblasts on the CaP coating showed enhanced attach- ment compared with the bare titanium surface. SEM obser- vations revealed osteoblast adherence on the coated surfaces (Fig. 6A and B). Cell survival and adhesion were also significantly enhanced with CaP-coated titanium (Fig. 6C and D): after 8 days of culture the survival rate was 70–80%, compared with 1% after 4 days culture on uncoated titanium alloy (Fig. 6D). mponent solutions (see Table 1); (D) four-component solutions; (E) SBF; Fig. 3. SEM, EDX and XRD analysis of a struvite-coated sample (A, C, and E) before and (B, D, and F) after additional coating with calcium phosphate solution. (�) indicates struvite peaks diffraction peaks; (+) indicates hydroxyapatite peaks. Fig. 2. SEM micrographs of Mg-sputtered titanium sheets after immersion in ADP solutions for periods of: (A) 30 s; (B) 2 min; (C) 15 min; and (D) 2 h. 2342 S. Ibasco et al. / Acta Biomaterialia 5 (2009) 2338–2347 S. Ibasco et al. / Acta Biomaterialia 5 (2009) 2338–2347 2343 4. Discussion In the present study we investigated the use of a reac- tive magnesium metal coating as both a nucleation site and a source of reactant cations for the formation of an apatitic layer on a titanium metal surface using a low-temperature aqueous deposition technique. Bioce- ramic coatings were produced by immersing magnesium sputter-coated titanium substrates in ADP and CaP solu- tions through chemical reactions between the magnesium metal and the ionic components of the solutions. We demonstrated an approach to produce reactant stable struvite coatings that could subsequently be replaced