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Electrochemically designed interfaces: Hydroxyapatite coated macro-mesoporous titania surfaces
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Electrochemically designed interfaces: Hydroxyapatite coated macro-mesoporous titania surfaces F.S. Utku a,∗ , E. Seckin b , G. Goller b , C. Tamerler c , M. Urgen b a b c
Yeditepe University, Department of Biomedical Engineering, Istanbul, Turkey Department of Metallurgical & Materials Engineering, Istanbul Technical University, Istanbul 06800, Turkey Department of Mechanical Engineering and Bioengineering Research Center, University of Kansas, Lawrence, KS 66045, USA
a r t i c l e
i n f o
Article history: Received 31 October 2014 Received in revised form 18 April 2015 Accepted 18 April 2015 Available online xxx Keywords: Macro-mesoporous titania Pulsed electrochemical deposition Hydroxyapatite coating Dual-acid polishing Surface functionalization
a b s t r a c t Titanium-based implants are key weight-bearing materials in biomedical engineering due to their excellent bulk mechanical properties and biocompatibility. Designing tissue-material interfaces of titanium implants is essential for an increase in osteointegration of engineered implant materials. Surface morphology is a crucial determinant in the construction of biocompatible and osteointegrative orthopedic and dental implants. Biomimicry of the structural features of bone, specifically its macro-to-mesoporosity, may enable the bone cells to osteointegrate, attain and maintain a physiological strain level. In this study, the surface chemistry and morphology of commercially pure titanium plates were modified using electrochemistry. Titanium oxide substrates were prepared by dual acid polishing and alkaline anodization using 0.1 M KOH in an electrochemical cell with a stainless steel cathode and an anodic voltage of 40 V at 20 ◦ C for 3 min. FE-SEM characterization revealed macro-mesoporous anodized titania surfaces, which were coated by hydroxyapatite using simulated body fluid and pulsed electrochemical deposition at 80 ◦ C, while unprocessed commercially pure titanium surfaces were used as controls. The calcium phosphate deposit on titania plates was characterized as calcium-deficient carbonated hydroxyapatite using XRD, FTIR and FE-SEM, whereas the deposit on non-porous, non-functionalized titanium surfaces was characterized as carbonated apatite. The adhesion strength of the hydroxyapatite coated titania surfaces was 38 ± 10 MPa, implying that these surfaces may be suitable for biological and chemical functionalization of medical implants to tune bioactivity, including delivering drugs. © 2015 Elsevier B.V. All rights reserved.
1. Introduction Titanium and alloys are widely used in the production of biomedical materials due to their high toughness, strength, elastic modulus, corrosion resistance, inertness and biocompatibility [1–3]. Titanium implant surfaces can be modified by physical deposition, thermo-chemical surface treatment, porous surface generation, ceramic coating in order to improve functional properties and osseointegration [2,4–7]. Complemented with the ease of surface microfunctionalization, titanium is used as weight-bearing orthopedic and dental material in the enhancement of osteointegration [8,9].
∗ Corresponding author at: Yeditepe University, Faculty of Engineering, Department of Biomedical Engineering, 26 Agustos Yerlesimi, Inonu Mahallesi, Kayisdagi Caddesi, 34755 Atasehir, Istanbul, Turkey. Tel.: +90 216 5780433; fax: +90 216 5780244. E-mail address: [email protected] (F.S. Utku).
Successful osteointegration of an implant, the ability of host tissues to interact with the implant without a layer of connective tissue, is a function of its surface topography, morphology, composition, chemistry and roughness [9–14]. With osteointegration, the implant displays increased mechanical stability, biological activity and chemical bonding [11,12]. Surface topography and substrate stiffness affect tissue-implant surface mechanical compatibility by the anisotropic stresses developed on the tissue [15] and dictate cellular attachment [16,17]. According to their phenotype and the nature of the cell adhesion receptor [18], mammalian cells display different cell morphologies and migration patterns [19] based on the wettability, roughness [20] and stiffness of the substrate [18]. Titanium implant surface can be modified through mechanical, physical and chemical means. Implants can be mechanically produced to have various shapes and surface topography by micro/machining, sand/grit blasting and rough polishing [9]. The physical methods of implant surface modification include atmospheric plasma and vacuum plasma spraying of Ca-P and TiO2 coatings, sputtering of thin films and ion deposition of titanium
http://dx.doi.org/10.1016/j.apsusc.2015.04.131 0169-4332/© 2015 Elsevier B.V. All rights reserved.
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implants. Chemical surface modification of titanium alters surface roughness and composition and enhances wettability of the implant [21]. The chemical methods used are hydrogen peroxide treatment, sol–gel treatment, chemical vapor deposition, acidic or alkaline treatment and anodization. The acid treatment used in this study entails the use of strong acids in order to remove the surface oxide and clean and polish the implant surface, giving a uniform roughness of 0.5–2 m and increasing not only the surface area, but also the migration and adhesion of osteogenic cells on the implant surface [22–26]. Alkaline treatment involves the use of sodium or potassium hydroxide and formation of a bioactive, rough TiO2 surface layer with higher biocompatibility, increased cell attachment and proliferation [27]. Anodization is an electrochemical process conducted in acidic or alkaline electrolyte solutions, where oxide films with improved adhesion and bonding are deposited on the titanium implant surface connected to the anodic electrode of an electrochemical cell. The oxide layer thickness can be modulated by altering the parameters of the electrochemical process, such as, current density, electrolyte solution concentration, composition and temperature. Produced by anodization in fluoride-containing electrolytes, nanotubular titania surfaces, provide a canaliculi-like tubular base for osteoblastic adhesion, leading to an advanced stage of cellular development [13,14,23,28]. The bonding strength of the titania nanotubes to the titanium base may pose a risk factor in biomedical applications, and therefore, must be either improved by annealing at high temperatures [29,30], or replaced by titania surfaces that can dissipate the bonding stress, i.e. macro-mesoporous surfaces with a higher bonding strength [30–32]. Production of a hierarchically organized macro-mesoporosity may overcome these mechanical concerns related to interfacial interactions at the implant-host level [33–35]. Mimicking the natural features of bone and using electrochemical deposition methods, ceramic coated biomaterials with macro- and/or mesoporous topography have been previously developed using acidic/(dual-acidic) and alkaline treatments [9,31]. The aim of this study is to produce a hierarchically organized macro-mesoporous implant surface using nitric-hydrofluoric acid, dual-acidic polishing, alkaline anodization and cathodic ceramic deposition, which, to the extent of our knowledge has not been previously experimented and implemented before. Here, a hierarchically organized topography, with mesoporosity embedded within the macroporosity, was produced at low temperatures using (nitric-hydrofluoric) dual-acid polishing and alkaline electrochemical anodization procedure and cathodically ceramic coated. 2. Materials and method 2.1. Materials Twelve commercially available pure titanium Grade IV (250 mm × 500 mm × 1 mm) plates were metallographically ground and polished using #120 grit down to #1200 grit Emery paper and finally 1 m diamond paste, washed with distilled water and sonicated in acetone. 2.2. Method 2.2.1. Surface functionalization Unprocessed titanium plates were used as controls. Titanium plates to be used as samples were dipped in 45%HNO3 :%20HF (v/v) aqueous solution for 10 min at room temperature and then rinsed in distilled water [8]. Consequently, titanium samples were anodized in 0.1 M KOH, in an electrochemical cell, using a stainless steel cathode electrode at an anodic voltage of 40 V at 20 ◦ C for 3 min (Fig. 1a).
Fig. 1. Schematic representation of titanium surface preparation using (a) acidic polishing and alkaline anodization and (b) cathodization methods. Acidic polishing of titanium produces macroporosity on the surface, while alkaline anodization introduces mesoporosity within the macroporous surface.
The surface morphology of the specimen was characterized using field emission scanning electron microscopy (FE-SEM) (JEOL JSM 7000F FEI). 2.2.2. Calcium phosphate coating All titanium plates were bioactivated by immersion in 0.5 M NaOH for 2 min at 50 ◦ C followed by rinsing in deionized water [36] prior to electrochemical deposition, which allows Ca-P coating of implants of any size and shape from an electrolyte solution [8,29]. Since the chemistry of calcium phosphate coating can be modulated by controlling certain parameters, such as current density, potential, length of cathodic and anodic deposition cycles, temperature and the type of solution, in this study, parameters were adjusted to enable deposition of hydroxyapatite [8,30,37–39]. Ca-P deposition was conducted using a modified simulated body fluid (SBF), containing reagent grade 0.15 M NaCl, 1.67 mM K2 HPO4 and 2.50 mM CaCl2 , pH buffered at 7.2 with the addition of 0.05 M tris(hydroxyl aminomethane) (pH 7.4) and hydrochloric acid [29]. A typical three-electrode electrochemical cell, with the titanium substrate as the working electrode, a standard silver–silver chloride electrode (Ag/AgCl in saturated KCl) as the reference electrode and the platinized stainless steel as the counter electrode, was used. At 80 ◦ C, surfaces were coated by pulsing the current density between −10 mA/cm2 for 0.2 s and 10 A/cm2 for 10 s for 100 pulsing cycles and then by maintaining a constant cathodic current density of −10 mA/cm2 for 60 min (Fig. 1b) [29]. 2.2.3. Characterization of calcium phosphate coating Samples were oven-dried at 100 ◦ C for 1 h and characterized using XRD, FT-IR and FE-SEM. A glancing X-ray diffractometer (Philips PW 3710), at the grazing angle of 1◦ , and at the range of scattering angle from 20 up to 80 2, at a step size of 0.2 2 was used to determine the XRD spectrum. FT-IR spectrometer (Perkin Elmer, Spectrum One) with the attenuated total reflectance (ATR) technique over a frequency range of 400–4000 cm−1 at a resolution of 4 cm−1 was used to examine molecular bonding. Sample crosssections were prepared by breaking the samples after immersion in liquid nitrogen for 15 s. The calcium and phosphorus ratios, surface and cross-sectional images of the coated titania plates were determined by field emission scanning electron microscope with an EDS spectrometer addition (FE-SEM) (JEOL JSM 7000F FEI). As a further analysis, the calcium and phosphorus content of the samples was determined respectively by atomic absorption spectrophotometer (AAS) (Perkin Elmer, Analyst 800) and by colorimetric
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3. Results 3.1. Surface characterization A micron–submicron scale, i.e. macro-mesoporous, oxide layer was formed on titanium samples (Fig. 2a). The FE-SEM images indicated formation of two types of surface features: wider and non-uniformly shaped micron to sub-micron macroporosity, ranging between 0.050 and 1.0 m in size, and inhomogenously distributed mesopores, below 0.050 m in size, within the macroporosity (Fig. 2b). 3.2. Characterization of Ca-P deposits
Fig. 2. (a) FE-SEM image of the acid polished and alkaline anodized titania plate, showing the macroporous surface structure. (Insert) A detail from the larger image showing the mesopores within the macroporosity. (b) FE-SEM image of the sample magnified at 320k times, showing the measured sizes of mesopores.
analysis at a wavelength of 690 nm using UV–vis spectrophotometer (Shimadzu, UV mini 240). Coated samples were mechanically tested at a rate of 1 mm/min according to ASTM 633-01, “Standard Test Method for Adhesion or Cohesion Strength of Thermal Spray Coatings” using a universal mechanical testing device (Instron, 1100).
The Ca-P layer deposited on samples was made of uniformly distributed acicular crystals of approximately 150 nm in diameter, 0.5 m in length (Fig. 3a) with an average Ca/P ratio of 1.43 ± 0.15, implying a reduction in calcium content and hydrogen phosphate incorporation in the lattice. Although the HA deposit demonstrated some interconnected porosity under the SEM, the samples were not studied to indicate formation of a mesoporous HA. The cross-sectional thickness of Ca-P films ranged between 3.0 and 5.0 m (Fig. 4a), with an average adhesion strength of 38 ± 10 MPa. The FE-SEM image of the coating after the mechanical testing is displayed in Fig. 4b. The 20–35◦ 2 range of the XRD spectrum displayed distinctive HA peaks at 25.8◦ (0 0 2), 28.1◦ (1 0 2), 28.97◦ (2 1 0), 31.8◦ (2 1 1), 32.1◦ (1 1 2), 32.9◦ (3 0 0) and 34.05◦ (2 0 2) 2, with an additional peak of 80% intensity appearing at 32.1◦ 2 (Fig. 5(i)) (ICSD 01-89-6437 for hydroxyapatite and ICSD 00-019-0272 for hydroxyapatite with a higher carbonate moiety). The FT-IR spectra of the deposits displayed the distinct peaks for HA. The phosphate peaks appeared at 560, 602, 961, 976–1190 cm−1 , weak hydroxide peaks appeared at 634 cm−1 for OH− libration and at 3574 cm−1 for OH stretch. The carbonate substitution for phosphate appeared as an additional peak at 875 cm−1 and at 1419 and 1454 cm−1 [40–44] (Fig. 6a). The cross-sectional thickness of Ca-P films deposited on control plates ranged between 3.3 and 4.0 m (Fig. 7). The deposit could be rubbed off by hand and thus control plates were not mechanically tested. The Ca-P layer was made of uniformly distributed acicular crystals of less than 100 nm in diameter and less than approximately 0.5 m in length (Fig. 3b), with Ca/P ratios of 1.44 ± 0.08. As was observed in the XRD pattern of the samples, the 32.11◦ 2 peak was expressed at 80% intensity in the controls (Fig. 5ii). The FTIR spectrum displayed splitting of the 1000–1200 cm−1 phosphate band into two major peaks at 1024 and 1117 cm−1 , implying the presence of hydrogen phosphates [45]. Controls displayed weaker
Fig. 3. FE-SEM image of (a) the CDHA coated porous sample surface and (b) carbonated apatite coated non-porous control surface.
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Fig. 4. FE-SEM image of (a) the cross-sectional view of the CDHA coated sample surface and (b) surface view of the sample after shear testing.
carbonate bands at 1430 and 1484 cm−1 and the hydroxide peaks at 625 and 3570 cm−1 were absent (Fig. 6b–d). 4. Discussion In this study, using dual-acid polishing and alkaline anodization methods, hierarchically arranged macro-mesoporous surfaces for enhanced osteointegrative implant-tissue interfaces could be efficiently produced on cp Grade IV titanium plates. Calcium-deficient carbonated hydroxyapatite (CDHA) and carbonated apatite deposition respectively on the functionalized macro-mesoporous and the unprocessed control substrates showed that the surface properties dictated the chemistry of the coating. 4.1. Formation of micro-mesoporous surface structure Acid polishing is a routine process used to remove debris and oxide scales in order to obtain clean and uniform surface finishes. Acid concentration, temperature and treatment time affect the extent of polishing, producing a macroporous surface. While titanium is oxidized to a bivalent state on the Ti and TiO2 interface, on the titania surface, oxygen ions are generated, which react with the Ti2+ cations to give TiO2 [8,9]. The concurrent reactions taking place at various interfaces of titania generate a TiO2 layer, which reacts with HF to form soluble titanium fluorides. Acid-etching and dual-acid polishing have been shown to increase osteointegration [46–49]. As the extent of metal surface cleaning/polishing depends on the acid concentration, temperature and treatment time, different mixtures of acids e.g. concentrated HNO3 and HF or concentrated HCl and H2 SO4 , heated above 100 ◦ C
Fig. 5. The XRD patterns of (i) the CDHA and (ii) the carbonated apatite deposits.
can be employed [9,23,46]. Using the dual acid-etching method, the attachment of fibrin and osteogenic cells, ECM production and direct bone formation on implant surface [24], adhesion of platelets, thus colonization of osteoblasts at the site and promotion of osseointegration with less bone resorption have been observed [47–49]. Anodization with alkaline solutions, introducing mesoporosity within the macropores, has also been shown to demonstrate osteoblastic colonization and osteointegration [8,9]. 4.2. Chemistry of the deposits Alkalinity and the ionic strength of solution, substrate surface, the surface control and diffusion control mechanisms, temperature-controlled reaction rates and kinetics of adsorption and diffusion govern the chemistry, particle size and morphology of the Ca-P compound deposited [50–52]. Using PECD, the chemistry of the Ca-P deposits is governed by (a) the reduction of water to hydroxide ions and thus conversion of acid phosphates to phosphates via alkalization at the cathode and/or electron donation, and (b) the diffusion of orthophosphate ions into the porous structure [8]. The type of orthophosphate ion reacting with calcium thus determines the type of Ca-P compound formed. These factors also govern the hydrogen phosphate content in solution [8,30,40–44], which may be substituted for phosphate ions, resulting in a decrease in anionic charge without a decrease in ionic content. Cationic and anionic co-substitutions then take place to neutralize the mineral; forming calcium deficient carbonated hydroxyapatite (CDHA) [40–44]. In this study, the samples and controls both had average Ca/P ratios of 1.4, and displayed the expected FT-IR phosphate, carbonate and hydrogen phosphate peaks, which indicate a decrease in anionic charge and thus a deficiency in calcium content in the mineral [45]. The presence of hydrogen phosphates in the compound may be due to partial conversion of hydrogen phosphates to phosphates in the neutral SBF solution, efficiently buffered by TRIS. The weak 3574 cm−1 hydroxide peak of the sample indicates formation of a compound similar to biological hydroxyapatite, specifically, a calcium deficient hydroxyapatite with hydrogen phosphate and carbonate substitutions for phosphate, as supported by the XRD data (Figs. 5 and 6). The presence of the 32.11◦ 2 peak at 80% intensity in the controls (as also is in the XRD data of the samples) may imply a carbonate moiety comparable to bone apatite [41–44] The carbonate substitution for phosphate appears as an additional peak at 875 cm−1 , indicating the presence of carbonate groups in the lattice of phosphate ions, and at 1419 and 1454 cm−1 , indicating the presence of surface carbonate ions interacting with water in the compound [40–44] (Fig. 6a). Weaker carbonate bands at 1430 and 1484 cm−1 indicate incorporation of carbonate in an apatitic compound (Fig. 6a). The absence of the 625, 3570 cm−1 hydroxide
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Fig. 6. The FT-IR spectra of (i) the CDHA and (ii) the carbonated apatite deposits. (a) The full spectra, (b) the 3562–3682 cm−1 interval for interval for the phosphate band, and (d) the 530–640 cm−1 interval for phosphate peak and OH libration.
peaks and the presence of a H O H stretch in the FT-IR spectrum (Fig. 6b and d) indicate formation of a less crystalline compound, which may be characterized as a carbonated apatite. 4.3. Mechanical properties of Ca-P deposits The success of bio-implantation is related to the chemistry of the surface coating, i.e. its surface energy, grain boundaries, chemistry, stoichiometry of the surface ions and crystallinity, all of which modulate the adsorption of organic biomolecules and ions in the tissue fluid. The strong ionic interactions in the highly crystalline deposits versus strong hydrogen bonding in less-crystalline compounds both contribute to the mechanical properties and the adhesion strength of the coating. In the literature, mainly the bonding strength has been studied [29], however, here, due to the torque forces that are applied on the implants, the unannealed shear strength has been studied. The bonding and shear strength of bone being ∼18 MPa and ∼44 MPa respectively, the high average unannealed shear strength of the deposits on the sample may be due to the presence of two interfaces, i.e. mineral–mineral (enabling chemical fixation) and mineral-macro-mesoporous titania (enabling biological
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OH stretch, (c) the 850–1200 cm−1
fixation). The FE-SEM images of mechanically tested samples (Fig. 4) indicated that despite the cracks formed, the coating was not completely stripped off the titanium base. The mechanical strength of the unannealed porous titania surfaces, was expected to further increase with annealing [29]. Surface roughness influences the behavior of osteoblastic cells, i.e. spreading and proliferation, differentiation, and protein synthesis [53]. While surface roughness of mm to 10 m scale contributes to primary long-term mechanical stability [54], roughness of 10–1 m contributes to biological fixation of the implant surface to the bone [54] and roughness of 1–100 nm scale enables adsorption of proteins and adhesion of osteoblastic cells [55] with various molecular processes taking place at the implant-tissue interface [56–60] Cellular adhesion to titania has been demonstrated to be a function of surface roughness, where the cells adhere to the surface through integrin receptors, causing changes in the cytoskeleton and thus leading to new gene expression [61]. Studies on this matter suggest that the substrate based conformational changes in cell shape affect membrane fluidity and calcium ion channels, altering gene expression and leading to attainment of a more advanced cellular development [22,56,61]. Osteoblasts, through nanometer to micron range (0.14–1.15 m) interactions with their environment,
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Fig. 7. FE-SEM image of the cross-sectional view of the carbonated apatite coating on the non-porous titania.
have been shown to attain osteointegration [22,56,61]. Therefore, although the ideal porosity for the bone making cells of 10–50 m in size may still be a debated question, it may be quite plausible to think that a hierarchically organized porosity attained by acidpolishing, producing an oxide layer as thin as 10 nm [8,9,30] and a surface roughness of 0.1 m to several m [8,9] may thus provide an adequate lacuna-like porous surface for the bone cells. 5. Conclusions The anodic and cathodic electrochemical processes have been used to produce macro-mesoporous, HA coated titania surfaces, with a shear strength, close to that of human bone. The porous implant surfaces are expected to biomimick bone surfaces and enhance osteointegration through the provision and improvement of the fundamental mechanical, biological and chemical elements of fixation. The macro-mesoporous surfaces thus may be potential implant-ceramic hard tissue interfaces for future biomedical applications. Acknowledgements This study has been supported by TUBITAK BIDEB 2218 PostDoctoral Research Project and TR-SPO. We thank Huseyin Sezer, Talat Apak, Berk Alkan, Sevgin Turkeli, Inci Kol, Muhammet Aydin and Mizrap Canibeyaz for their help in the characterization of titania substrates and HA coatings. References [1] M. Long, H.J. Rack, Titanium alloys in total joint replacement – a materials science perspective, Biomaterials 19 (1998) 1621–1639. [2] M. Geetha, A.K. Singh, R. Asokamani, A.K. Gogia, Ti based biomaterials, the ultimate choice for orthopaedic implants – a review, Prog. Mater. Sci. 54 (2009) 397–425. [3] X.Y. Liu, P.K. Chu, C.X. Ding, Surface modification of titanium, titanium alloys, and related materials for biomedical applications, Mater. Sci. Eng. R: Rep. 47 (2004) 49–121. [4] S. Fukumoto, H. Tsubakino, S. Inoue, L. Liu, M. Terasawa, T. Mitamura, Surface modification of titanium by nitrogen ion implantation, Mater. Sci. Eng. A: Struct. Mater. Prop. Microstruct. Proc. 263 (1999) 205–209. [5] C. Alves, C. Neto, G.H.S. Morais, C.F. Da Silva, V. Hajek, Nitriding of titanium disks and industrial dental implants using hollow cathode discharge, Surf. Coat. Technol. 194 (2005) 196–202. [6] D. Stojanovic, B. Jokic, D. Veljovic, R. Petrovic, P.S. Uskokovic, D. Janackovic, Bioactive glass–apatite composite coating for titanium implant synthesized by electrophoretic deposition, J. Eur. Ceram. Soc. 27 (2007) 1595–1599.
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Electrochemically designed interfaces: Hydroxyapatite coated macro-mesoporous titania surfaces F.S. Utku a,∗ , E. Seckin b , G. Goller b , C. Tamerler c , M. Urgen b a b c
Yeditepe University, Department of Biomedical Engineering, Istanbul, Turkey Department of Metallurgical & Materials Engineering, Istanbul Technical University, Istanbul 06800, Turkey Department of Mechanical Engineering and Bioengineering Research Center, University of Kansas, Lawrence, KS 66045, USA
a r t i c l e
i n f o
Article history: Received 31 October 2014 Received in revised form 18 April 2015 Accepted 18 April 2015 Available online xxx Keywords: Macro-mesoporous titania Pulsed electrochemical deposition Hydroxyapatite coating Dual-acid polishing Surface functionalization
a b s t r a c t Titanium-based implants are key weight-bearing materials in biomedical engineering due to their excellent bulk mechanical properties and biocompatibility. Designing tissue-material interfaces of titanium implants is essential for an increase in osteointegration of engineered implant materials. Surface morphology is a crucial determinant in the construction of biocompatible and osteointegrative orthopedic and dental implants. Biomimicry of the structural features of bone, specifically its macro-to-mesoporosity, may enable the bone cells to osteointegrate, attain and maintain a physiological strain level. In this study, the surface chemistry and morphology of commercially pure titanium plates were modified using electrochemistry. Titanium oxide substrates were prepared by dual acid polishing and alkaline anodization using 0.1 M KOH in an electrochemical cell with a stainless steel cathode and an anodic voltage of 40 V at 20 ◦ C for 3 min. FE-SEM characterization revealed macro-mesoporous anodized titania surfaces, which were coated by hydroxyapatite using simulated body fluid and pulsed electrochemical deposition at 80 ◦ C, while unprocessed commercially pure titanium surfaces were used as controls. The calcium phosphate deposit on titania plates was characterized as calcium-deficient carbonated hydroxyapatite using XRD, FTIR and FE-SEM, whereas the deposit on non-porous, non-functionalized titanium surfaces was characterized as carbonated apatite. The adhesion strength of the hydroxyapatite coated titania surfaces was 38 ± 10 MPa, implying that these surfaces may be suitable for biological and chemical functionalization of medical implants to tune bioactivity, including delivering drugs. © 2015 Elsevier B.V. All rights reserved.
1. Introduction Titanium and alloys are widely used in the production of biomedical materials due to their high toughness, strength, elastic modulus, corrosion resistance, inertness and biocompatibility [1–3]. Titanium implant surfaces can be modified by physical deposition, thermo-chemical surface treatment, porous surface generation, ceramic coating in order to improve functional properties and osseointegration [2,4–7]. Complemented with the ease of surface microfunctionalization, titanium is used as weight-bearing orthopedic and dental material in the enhancement of osteointegration [8,9].
∗ Corresponding author at: Yeditepe University, Faculty of Engineering, Department of Biomedical Engineering, 26 Agustos Yerlesimi, Inonu Mahallesi, Kayisdagi Caddesi, 34755 Atasehir, Istanbul, Turkey. Tel.: +90 216 5780433; fax: +90 216 5780244. E-mail address: [email protected] (F.S. Utku).
Successful osteointegration of an implant, the ability of host tissues to interact with the implant without a layer of connective tissue, is a function of its surface topography, morphology, composition, chemistry and roughness [9–14]. With osteointegration, the implant displays increased mechanical stability, biological activity and chemical bonding [11,12]. Surface topography and substrate stiffness affect tissue-implant surface mechanical compatibility by the anisotropic stresses developed on the tissue [15] and dictate cellular attachment [16,17]. According to their phenotype and the nature of the cell adhesion receptor [18], mammalian cells display different cell morphologies and migration patterns [19] based on the wettability, roughness [20] and stiffness of the substrate [18]. Titanium implant surface can be modified through mechanical, physical and chemical means. Implants can be mechanically produced to have various shapes and surface topography by micro/machining, sand/grit blasting and rough polishing [9]. The physical methods of implant surface modification include atmospheric plasma and vacuum plasma spraying of Ca-P and TiO2 coatings, sputtering of thin films and ion deposition of titanium
http://dx.doi.org/10.1016/j.apsusc.2015.04.131 0169-4332/© 2015 Elsevier B.V. All rights reserved.
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implants. Chemical surface modification of titanium alters surface roughness and composition and enhances wettability of the implant [21]. The chemical methods used are hydrogen peroxide treatment, sol–gel treatment, chemical vapor deposition, acidic or alkaline treatment and anodization. The acid treatment used in this study entails the use of strong acids in order to remove the surface oxide and clean and polish the implant surface, giving a uniform roughness of 0.5–2 m and increasing not only the surface area, but also the migration and adhesion of osteogenic cells on the implant surface [22–26]. Alkaline treatment involves the use of sodium or potassium hydroxide and formation of a bioactive, rough TiO2 surface layer with higher biocompatibility, increased cell attachment and proliferation [27]. Anodization is an electrochemical process conducted in acidic or alkaline electrolyte solutions, where oxide films with improved adhesion and bonding are deposited on the titanium implant surface connected to the anodic electrode of an electrochemical cell. The oxide layer thickness can be modulated by altering the parameters of the electrochemical process, such as, current density, electrolyte solution concentration, composition and temperature. Produced by anodization in fluoride-containing electrolytes, nanotubular titania surfaces, provide a canaliculi-like tubular base for osteoblastic adhesion, leading to an advanced stage of cellular development [13,14,23,28]. The bonding strength of the titania nanotubes to the titanium base may pose a risk factor in biomedical applications, and therefore, must be either improved by annealing at high temperatures [29,30], or replaced by titania surfaces that can dissipate the bonding stress, i.e. macro-mesoporous surfaces with a higher bonding strength [30–32]. Production of a hierarchically organized macro-mesoporosity may overcome these mechanical concerns related to interfacial interactions at the implant-host level [33–35]. Mimicking the natural features of bone and using electrochemical deposition methods, ceramic coated biomaterials with macro- and/or mesoporous topography have been previously developed using acidic/(dual-acidic) and alkaline treatments [9,31]. The aim of this study is to produce a hierarchically organized macro-mesoporous implant surface using nitric-hydrofluoric acid, dual-acidic polishing, alkaline anodization and cathodic ceramic deposition, which, to the extent of our knowledge has not been previously experimented and implemented before. Here, a hierarchically organized topography, with mesoporosity embedded within the macroporosity, was produced at low temperatures using (nitric-hydrofluoric) dual-acid polishing and alkaline electrochemical anodization procedure and cathodically ceramic coated. 2. Materials and method 2.1. Materials Twelve commercially available pure titanium Grade IV (250 mm × 500 mm × 1 mm) plates were metallographically ground and polished using #120 grit down to #1200 grit Emery paper and finally 1 m diamond paste, washed with distilled water and sonicated in acetone. 2.2. Method 2.2.1. Surface functionalization Unprocessed titanium plates were used as controls. Titanium plates to be used as samples were dipped in 45%HNO3 :%20HF (v/v) aqueous solution for 10 min at room temperature and then rinsed in distilled water [8]. Consequently, titanium samples were anodized in 0.1 M KOH, in an electrochemical cell, using a stainless steel cathode electrode at an anodic voltage of 40 V at 20 ◦ C for 3 min (Fig. 1a).
Fig. 1. Schematic representation of titanium surface preparation using (a) acidic polishing and alkaline anodization and (b) cathodization methods. Acidic polishing of titanium produces macroporosity on the surface, while alkaline anodization introduces mesoporosity within the macroporous surface.
The surface morphology of the specimen was characterized using field emission scanning electron microscopy (FE-SEM) (JEOL JSM 7000F FEI). 2.2.2. Calcium phosphate coating All titanium plates were bioactivated by immersion in 0.5 M NaOH for 2 min at 50 ◦ C followed by rinsing in deionized water [36] prior to electrochemical deposition, which allows Ca-P coating of implants of any size and shape from an electrolyte solution [8,29]. Since the chemistry of calcium phosphate coating can be modulated by controlling certain parameters, such as current density, potential, length of cathodic and anodic deposition cycles, temperature and the type of solution, in this study, parameters were adjusted to enable deposition of hydroxyapatite [8,30,37–39]. Ca-P deposition was conducted using a modified simulated body fluid (SBF), containing reagent grade 0.15 M NaCl, 1.67 mM K2 HPO4 and 2.50 mM CaCl2 , pH buffered at 7.2 with the addition of 0.05 M tris(hydroxyl aminomethane) (pH 7.4) and hydrochloric acid [29]. A typical three-electrode electrochemical cell, with the titanium substrate as the working electrode, a standard silver–silver chloride electrode (Ag/AgCl in saturated KCl) as the reference electrode and the platinized stainless steel as the counter electrode, was used. At 80 ◦ C, surfaces were coated by pulsing the current density between −10 mA/cm2 for 0.2 s and 10 A/cm2 for 10 s for 100 pulsing cycles and then by maintaining a constant cathodic current density of −10 mA/cm2 for 60 min (Fig. 1b) [29]. 2.2.3. Characterization of calcium phosphate coating Samples were oven-dried at 100 ◦ C for 1 h and characterized using XRD, FT-IR and FE-SEM. A glancing X-ray diffractometer (Philips PW 3710), at the grazing angle of 1◦ , and at the range of scattering angle from 20 up to 80 2, at a step size of 0.2 2 was used to determine the XRD spectrum. FT-IR spectrometer (Perkin Elmer, Spectrum One) with the attenuated total reflectance (ATR) technique over a frequency range of 400–4000 cm−1 at a resolution of 4 cm−1 was used to examine molecular bonding. Sample crosssections were prepared by breaking the samples after immersion in liquid nitrogen for 15 s. The calcium and phosphorus ratios, surface and cross-sectional images of the coated titania plates were determined by field emission scanning electron microscope with an EDS spectrometer addition (FE-SEM) (JEOL JSM 7000F FEI). As a further analysis, the calcium and phosphorus content of the samples was determined respectively by atomic absorption spectrophotometer (AAS) (Perkin Elmer, Analyst 800) and by colorimetric
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3. Results 3.1. Surface characterization A micron–submicron scale, i.e. macro-mesoporous, oxide layer was formed on titanium samples (Fig. 2a). The FE-SEM images indicated formation of two types of surface features: wider and non-uniformly shaped micron to sub-micron macroporosity, ranging between 0.050 and 1.0 m in size, and inhomogenously distributed mesopores, below 0.050 m in size, within the macroporosity (Fig. 2b). 3.2. Characterization of Ca-P deposits
Fig. 2. (a) FE-SEM image of the acid polished and alkaline anodized titania plate, showing the macroporous surface structure. (Insert) A detail from the larger image showing the mesopores within the macroporosity. (b) FE-SEM image of the sample magnified at 320k times, showing the measured sizes of mesopores.
analysis at a wavelength of 690 nm using UV–vis spectrophotometer (Shimadzu, UV mini 240). Coated samples were mechanically tested at a rate of 1 mm/min according to ASTM 633-01, “Standard Test Method for Adhesion or Cohesion Strength of Thermal Spray Coatings” using a universal mechanical testing device (Instron, 1100).
The Ca-P layer deposited on samples was made of uniformly distributed acicular crystals of approximately 150 nm in diameter, 0.5 m in length (Fig. 3a) with an average Ca/P ratio of 1.43 ± 0.15, implying a reduction in calcium content and hydrogen phosphate incorporation in the lattice. Although the HA deposit demonstrated some interconnected porosity under the SEM, the samples were not studied to indicate formation of a mesoporous HA. The cross-sectional thickness of Ca-P films ranged between 3.0 and 5.0 m (Fig. 4a), with an average adhesion strength of 38 ± 10 MPa. The FE-SEM image of the coating after the mechanical testing is displayed in Fig. 4b. The 20–35◦ 2 range of the XRD spectrum displayed distinctive HA peaks at 25.8◦ (0 0 2), 28.1◦ (1 0 2), 28.97◦ (2 1 0), 31.8◦ (2 1 1), 32.1◦ (1 1 2), 32.9◦ (3 0 0) and 34.05◦ (2 0 2) 2, with an additional peak of 80% intensity appearing at 32.1◦ 2 (Fig. 5(i)) (ICSD 01-89-6437 for hydroxyapatite and ICSD 00-019-0272 for hydroxyapatite with a higher carbonate moiety). The FT-IR spectra of the deposits displayed the distinct peaks for HA. The phosphate peaks appeared at 560, 602, 961, 976–1190 cm−1 , weak hydroxide peaks appeared at 634 cm−1 for OH− libration and at 3574 cm−1 for OH stretch. The carbonate substitution for phosphate appeared as an additional peak at 875 cm−1 and at 1419 and 1454 cm−1 [40–44] (Fig. 6a). The cross-sectional thickness of Ca-P films deposited on control plates ranged between 3.3 and 4.0 m (Fig. 7). The deposit could be rubbed off by hand and thus control plates were not mechanically tested. The Ca-P layer was made of uniformly distributed acicular crystals of less than 100 nm in diameter and less than approximately 0.5 m in length (Fig. 3b), with Ca/P ratios of 1.44 ± 0.08. As was observed in the XRD pattern of the samples, the 32.11◦ 2 peak was expressed at 80% intensity in the controls (Fig. 5ii). The FTIR spectrum displayed splitting of the 1000–1200 cm−1 phosphate band into two major peaks at 1024 and 1117 cm−1 , implying the presence of hydrogen phosphates [45]. Controls displayed weaker
Fig. 3. FE-SEM image of (a) the CDHA coated porous sample surface and (b) carbonated apatite coated non-porous control surface.
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Fig. 4. FE-SEM image of (a) the cross-sectional view of the CDHA coated sample surface and (b) surface view of the sample after shear testing.
carbonate bands at 1430 and 1484 cm−1 and the hydroxide peaks at 625 and 3570 cm−1 were absent (Fig. 6b–d). 4. Discussion In this study, using dual-acid polishing and alkaline anodization methods, hierarchically arranged macro-mesoporous surfaces for enhanced osteointegrative implant-tissue interfaces could be efficiently produced on cp Grade IV titanium plates. Calcium-deficient carbonated hydroxyapatite (CDHA) and carbonated apatite deposition respectively on the functionalized macro-mesoporous and the unprocessed control substrates showed that the surface properties dictated the chemistry of the coating. 4.1. Formation of micro-mesoporous surface structure Acid polishing is a routine process used to remove debris and oxide scales in order to obtain clean and uniform surface finishes. Acid concentration, temperature and treatment time affect the extent of polishing, producing a macroporous surface. While titanium is oxidized to a bivalent state on the Ti and TiO2 interface, on the titania surface, oxygen ions are generated, which react with the Ti2+ cations to give TiO2 [8,9]. The concurrent reactions taking place at various interfaces of titania generate a TiO2 layer, which reacts with HF to form soluble titanium fluorides. Acid-etching and dual-acid polishing have been shown to increase osteointegration [46–49]. As the extent of metal surface cleaning/polishing depends on the acid concentration, temperature and treatment time, different mixtures of acids e.g. concentrated HNO3 and HF or concentrated HCl and H2 SO4 , heated above 100 ◦ C
Fig. 5. The XRD patterns of (i) the CDHA and (ii) the carbonated apatite deposits.
can be employed [9,23,46]. Using the dual acid-etching method, the attachment of fibrin and osteogenic cells, ECM production and direct bone formation on implant surface [24], adhesion of platelets, thus colonization of osteoblasts at the site and promotion of osseointegration with less bone resorption have been observed [47–49]. Anodization with alkaline solutions, introducing mesoporosity within the macropores, has also been shown to demonstrate osteoblastic colonization and osteointegration [8,9]. 4.2. Chemistry of the deposits Alkalinity and the ionic strength of solution, substrate surface, the surface control and diffusion control mechanisms, temperature-controlled reaction rates and kinetics of adsorption and diffusion govern the chemistry, particle size and morphology of the Ca-P compound deposited [50–52]. Using PECD, the chemistry of the Ca-P deposits is governed by (a) the reduction of water to hydroxide ions and thus conversion of acid phosphates to phosphates via alkalization at the cathode and/or electron donation, and (b) the diffusion of orthophosphate ions into the porous structure [8]. The type of orthophosphate ion reacting with calcium thus determines the type of Ca-P compound formed. These factors also govern the hydrogen phosphate content in solution [8,30,40–44], which may be substituted for phosphate ions, resulting in a decrease in anionic charge without a decrease in ionic content. Cationic and anionic co-substitutions then take place to neutralize the mineral; forming calcium deficient carbonated hydroxyapatite (CDHA) [40–44]. In this study, the samples and controls both had average Ca/P ratios of 1.4, and displayed the expected FT-IR phosphate, carbonate and hydrogen phosphate peaks, which indicate a decrease in anionic charge and thus a deficiency in calcium content in the mineral [45]. The presence of hydrogen phosphates in the compound may be due to partial conversion of hydrogen phosphates to phosphates in the neutral SBF solution, efficiently buffered by TRIS. The weak 3574 cm−1 hydroxide peak of the sample indicates formation of a compound similar to biological hydroxyapatite, specifically, a calcium deficient hydroxyapatite with hydrogen phosphate and carbonate substitutions for phosphate, as supported by the XRD data (Figs. 5 and 6). The presence of the 32.11◦ 2 peak at 80% intensity in the controls (as also is in the XRD data of the samples) may imply a carbonate moiety comparable to bone apatite [41–44] The carbonate substitution for phosphate appears as an additional peak at 875 cm−1 , indicating the presence of carbonate groups in the lattice of phosphate ions, and at 1419 and 1454 cm−1 , indicating the presence of surface carbonate ions interacting with water in the compound [40–44] (Fig. 6a). Weaker carbonate bands at 1430 and 1484 cm−1 indicate incorporation of carbonate in an apatitic compound (Fig. 6a). The absence of the 625, 3570 cm−1 hydroxide
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Fig. 6. The FT-IR spectra of (i) the CDHA and (ii) the carbonated apatite deposits. (a) The full spectra, (b) the 3562–3682 cm−1 interval for interval for the phosphate band, and (d) the 530–640 cm−1 interval for phosphate peak and OH libration.
peaks and the presence of a H O H stretch in the FT-IR spectrum (Fig. 6b and d) indicate formation of a less crystalline compound, which may be characterized as a carbonated apatite. 4.3. Mechanical properties of Ca-P deposits The success of bio-implantation is related to the chemistry of the surface coating, i.e. its surface energy, grain boundaries, chemistry, stoichiometry of the surface ions and crystallinity, all of which modulate the adsorption of organic biomolecules and ions in the tissue fluid. The strong ionic interactions in the highly crystalline deposits versus strong hydrogen bonding in less-crystalline compounds both contribute to the mechanical properties and the adhesion strength of the coating. In the literature, mainly the bonding strength has been studied [29], however, here, due to the torque forces that are applied on the implants, the unannealed shear strength has been studied. The bonding and shear strength of bone being ∼18 MPa and ∼44 MPa respectively, the high average unannealed shear strength of the deposits on the sample may be due to the presence of two interfaces, i.e. mineral–mineral (enabling chemical fixation) and mineral-macro-mesoporous titania (enabling biological
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OH stretch, (c) the 850–1200 cm−1
fixation). The FE-SEM images of mechanically tested samples (Fig. 4) indicated that despite the cracks formed, the coating was not completely stripped off the titanium base. The mechanical strength of the unannealed porous titania surfaces, was expected to further increase with annealing [29]. Surface roughness influences the behavior of osteoblastic cells, i.e. spreading and proliferation, differentiation, and protein synthesis [53]. While surface roughness of mm to 10 m scale contributes to primary long-term mechanical stability [54], roughness of 10–1 m contributes to biological fixation of the implant surface to the bone [54] and roughness of 1–100 nm scale enables adsorption of proteins and adhesion of osteoblastic cells [55] with various molecular processes taking place at the implant-tissue interface [56–60] Cellular adhesion to titania has been demonstrated to be a function of surface roughness, where the cells adhere to the surface through integrin receptors, causing changes in the cytoskeleton and thus leading to new gene expression [61]. Studies on this matter suggest that the substrate based conformational changes in cell shape affect membrane fluidity and calcium ion channels, altering gene expression and leading to attainment of a more advanced cellular development [22,56,61]. Osteoblasts, through nanometer to micron range (0.14–1.15 m) interactions with their environment,
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Fig. 7. FE-SEM image of the cross-sectional view of the carbonated apatite coating on the non-porous titania.
have been shown to attain osteointegration [22,56,61]. Therefore, although the ideal porosity for the bone making cells of 10–50 m in size may still be a debated question, it may be quite plausible to think that a hierarchically organized porosity attained by acidpolishing, producing an oxide layer as thin as 10 nm [8,9,30] and a surface roughness of 0.1 m to several m [8,9] may thus provide an adequate lacuna-like porous surface for the bone cells. 5. Conclusions The anodic and cathodic electrochemical processes have been used to produce macro-mesoporous, HA coated titania surfaces, with a shear strength, close to that of human bone. The porous implant surfaces are expected to biomimick bone surfaces and enhance osteointegration through the provision and improvement of the fundamental mechanical, biological and chemical elements of fixation. The macro-mesoporous surfaces thus may be potential implant-ceramic hard tissue interfaces for future biomedical applications. Acknowledgements This study has been supported by TUBITAK BIDEB 2218 PostDoctoral Research Project and TR-SPO. We thank Huseyin Sezer, Talat Apak, Berk Alkan, Sevgin Turkeli, Inci Kol, Muhammet Aydin and Mizrap Canibeyaz for their help in the characterization of titania substrates and HA coatings. References [1] M. Long, H.J. Rack, Titanium alloys in total joint replacement – a materials science perspective, Biomaterials 19 (1998) 1621–1639. [2] M. Geetha, A.K. Singh, R. Asokamani, A.K. Gogia, Ti based biomaterials, the ultimate choice for orthopaedic implants – a review, Prog. Mater. Sci. 54 (2009) 397–425. [3] X.Y. Liu, P.K. Chu, C.X. Ding, Surface modification of titanium, titanium alloys, and related materials for biomedical applications, Mater. Sci. Eng. R: Rep. 47 (2004) 49–121. [4] S. Fukumoto, H. Tsubakino, S. Inoue, L. Liu, M. Terasawa, T. Mitamura, Surface modification of titanium by nitrogen ion implantation, Mater. Sci. Eng. A: Struct. Mater. Prop. Microstruct. Proc. 263 (1999) 205–209. [5] C. Alves, C. Neto, G.H.S. Morais, C.F. Da Silva, V. Hajek, Nitriding of titanium disks and industrial dental implants using hollow cathode discharge, Surf. Coat. Technol. 194 (2005) 196–202. [6] D. Stojanovic, B. Jokic, D. Veljovic, R. Petrovic, P.S. Uskokovic, D. Janackovic, Bioactive glass–apatite composite coating for titanium implant synthesized by electrophoretic deposition, J. Eur. Ceram. Soc. 27 (2007) 1595–1599.
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