Functional coatings formed on the titanium and magnesium alloys as implant materials by plasma electrolytic oxidation technology: fundamental principles and synthesis conditions

1 Introduction

Artificial materials are extensively used in advanced medicine for replacement of damaged tissues and organs. Depending on their purpose, implants incorporated into the organism should be further replaced by a living tissue and/or function during a long period. In general, materials application in implants should meet a number of requirements, in particular: they must not corrode, must have mechanical characteristics similar to those of the bone tissue, must not cause reactions of the immune system, must be integrated with the bone tissue, and must stimulate osteogenesis (Cao et al., 2015; Matusiewicz, 2014; Metoki et al., 2014a; Pound, 2014a; Wu et al., 2014).

The majority of metal implants presently used in implant surgery are titanium materials owing to their excellent corrosion and corrosion-durable properties and biocompatibility as well as good (but not perfect) mechanical features (Bolelli et al., 2014; Calderon-Moreno et al., 2014; Eliaz et al., 2011; Lakstein et al., 2009; Pound, 2014b). At the same time, bioactivity is the major drawback of Ti medical implants, e.g. commercially pure titanium, Ti-6Al-4V, Ti-6Al-7Nb (Campanelli, Duarte, da Silva, & Bolfarini, 2014; Duarte, Bolfarini, Biaggio, Rocha-Filho, & Nascente, 2014; Eliaz, 2012; Guo et al., 2015; Zieliński & Sobieszczyk, 2010). A promising method for increasing the strength of bone tissue joints with implants consists of deposition of coatings of bioactive materials with suitable compatibility with the organism. Because of the similarity in chemical composition, calcium phosphate-based materials find extensive application as bioactive ceramics for restoration of integrity of damaged bones (Arifin, Sulong, Muhamad, Syarif, & Ramli, 2014; Fathi, Hanifi, & Mortazavi, 2008; Han, Hong, & Xu, 2003; Sridhar, Eliaz, Mudali, & Raj, 2002). In particular, physical and chemical properties of hydroxyapatite (HA), Ca₁₀(PO₄)₆(OH)₂, provide acceptable biocompatibility actively stimulating osteogenesis and restoration of the bone tissue (Eliaz, 2010; Skogareva, Ivanov, Baranchikov, Minaeva, & Tripol’skaya, 2015). In accordance with Safronova, Putlyaev, Shekhirev, and Kuznetsov (2007), to increase biocompatibility with the bone tissue, non-stoichiometric HA (Ca_10-x(HPO₄)_x(PO₄)_6-x(OH)_2-x, 0<x<1) diluted with more soluble calcium phosphates, for example, triple-substituted orthophosphates, is sometimes used. However, HA is susceptible to fatigue failure, making it unsuitable for load-bearing materials (Arifin et al., 2014; Fathi et al., 2008; Han et al., 2003; Sridhar et al., 2002). Thus, to preserve the mechanical properties of the titanium, e.g. load-bearing ability, and at the same time take advantage of the surface’s chemical similarity, bioactivity, and biocompatibility with the bone, the formation of the HA-containing coatings on Ti substrates is required (Evdokimov et al., 2014; Skogareva, Ivanov, Pilipenko, & Tripol’skaya, 2012). Here, the bond strength between the HA layer and the metal surface is a critical factor. Peeling of the coating from the implant in the human body results in adverse effects on the implants and the surrounding tissue due to detached particles (Eliaz, Kopelovitch, Burstein, Kobayashi, & Hanawa, 2009a; Eliaz et al., 2009b; Kim, Jeong, Choe, & Brantley, 2013; Suchanek, Yashma, Kakihana, & Yoshimura, 1997). To increase the adhesive strength of coatings to the titanium substrate, in some cases, it is necessary, prior to their deposition, to treat the surface to create a certain roughness. The treatment can be sandblasting or chemical etching, which creates a multilevel (multimodal) surface with a very porous structure, which in turn increases the adhesive strength of the calcium phosphate coatings to the titanium substrate (Mohseni, Zalnezhad, & Bushroa, 2014; Zavgorodniy, Mason, LeGeros, & Rohanizadeh, 2012).

Unlike other Ti biomaterials, the most important problem for the NiTi alloy (46–52 at% Ti and 48–54 at% Ni), which is characterized by a unique shape memory effect for implant surgery, is the need for a protective coating on the surface (Ng, Man, & Yue, 2008; Rondelli, 1996; Sun & Wang, 2008). Diffusion and accumulation of Ni²⁺ ions in the soft tissues of the organism cause negative consequences, in particular, formation of neoplasms (Habijan et al., 2011; Tosun et al., 2013). To protect the implant against a corrosion-active biological environment and to provide better adaptation of bone tissues to the foreign body, the optimal variant will be the formation of hetero-oxide composite bioinert layers on the NiTi surface (Liu, Chu, & Ding, 2004; Xu et al., 2010; Xu, Liu, Luo, & Zhao, 2013).

Magnesium alloys have been increasingly getting more attention, as they have biodegradable implant applications. The major advantages of these materials include biocompatibility and suitable mechanical properties, such as low density and Young’s modulus comparable to those of cortical bone (Narayanan, Park, & Lee, 2015a,b; Raman, Jafari, & Harandi, 2015). According to reports, anodic dissolution of Mg yields Mg²⁺ cations, which are generally non-toxic and cause no adverse side effects (Narayanan et al., 2015b). Moreover, despite different opinions, hydrogen evolution is not the main problem of Mg implant materials (Narayanan et al., 2015a,b). In general, there are no enzymatic functions based on hydrogen reaction in the human body (Nakao, 2011). However, due to anti-oxidant and anti-inflammatory properties, H₂ has a favorable influence on cell and tissue protection from harmful effects of active oxygen and free radicals (Buchholz et al., 2008; Huang, Kawamura, Toyoda, & Nakao, 2010; Nakao, 2011). In addition, H₂ dissolution in blood leads to lowering of systolic arterial blood pressure (Buchholz et al., 2008; Huang et al., 2010; Nakao, 2011; Nakayama, Nakano, & Nakayama, 2010). The recent results confirmed that hydrogen neither has adverse reactions in osteointegration nor does it inhibit bone growth (Kraus et al., 2012). The principal issue associated with H₂ presence in organisms concerns subcutaneous gas bubbles, and dimensions of which depend on hydrogen concentration. Such bubbles cause necrosis of the neighboring tissues and impede healing (Olinger, 2000; Roewer, Thiel, & Wunder, 2012; Staiger, Pietak, Huadmai, & Dias, 2006). In this case, surgical H₂ removal, wherein gas is moved out of the body, is a key point in solving the problem of hydrogen in resorbable Mg implants. The important problems in using magnesium implants are poor cell adhesion, difficulty in the formation of the new bone tissue, and premature loss of the implant’s mechanical strength due to extremely high corrosion activity of Mg in chloride-containing environments (Narayanan et al., 2015a,b). In addition, some magnesium alloys include toxic alloying elements or impurities that could be released during degradation. Taking into consideration the requirements of biocompatibility, only a limited number of elements are suitable for Mg alloys used in medical applications, e.g. Mn, Zn, Sr, Ce, Ca (Carboneras, Garcia-Alonso, & Escudero, 2011; Groysman, 2010; Zeng, Dietzel, Witte, Hort, & Blawert, 2008). Magnesium corrosion rate can be reduced by the application of protective coatings. However, for resorbable Mg implants, it is essential that new bone tissue growth proceeds before implant dissolution (Hagihara, Shakudo, Fujii, & Nakano, 2014; Hiromoto et al., 2008; Maiorova, Safonov, & Skundin, 2013; Virtanen, 2010). Therefore, the protective coating must have a certain bioactivity, i.e. accelerate bone regeneration while inhibiting implant corrosion, thus promoting its gradual replacement with osteoblasts. Designing such multifunctional coatings that are both corrosion resistant and bioactive is vital for resorbable Mg alloy implants. In this context, bioactive coatings containing native to the bone tissue calcium phosphates are of great interest (Chen, Dai, & Zhang, 2015; Zhang et al., 2015).

Plasma electrolytic oxidation (PEO) is based on DC or AC polarization of a material treated at high voltage that produces plasma micro-discharges on the electrode surface, yielding high temperature (up to 5000–7000 K) and high pressure (up to 100 MPa) in its channels. As a result of local high-energy influence, layers, including both matrix (metal to be oxidized) and electrolyte elements, are formed on the product surface (Boinovich et al., 2012). The properties of these layers differ from those of conventional anode oxide films. Subsequent treatment of the surface produced by PEO (e.g. filling of pores with bioinert polymer) can be used to form composite coatings that have promising application in implant surgery (Malayoglu, Tekin, & Shrestha, 2010; Mohedano, Blawert, & Zheludkevich, 2015).

This review presents the results confirming that PEO is a suitable technology for surface modification of both titanium and magnesium alloys used in implant surgery. The fundamental principles of formation of bioactive calcium phosphate surface PEO layers on commercially pure titanium VT1-0, the bioinert protective composite coatings on NiTi, and the corrosion-resistant, bioactive PEO coatings containing HA on Mg-Mn-Ce magnesium alloy have been discussed in detail.

2 Bioactive calcium phosphate coatings

Titanium is a biomaterial extensively used in implants for orthopedic dental and orthodontic devices owing to its excellent corrosion resistance as well as mechanical strength (Al-Noaman, Karpukhina, Rawlinson, & Hill, 2013a,b; Ferraris, Bobbio, Miola, & Spriano, 2015; Li et al., 2015; Tsai et al., 2013). However, the bioinert character of Ti poses a problem for its application in implants. Surface modification by bioactive Ca- and P-containing anatase-based ceramic coatings is a promising way to improve the osteoinductive properties of titanium. In particular, Li et al. (2015) showed that osteointegration and the antibacterial properties of Ti implants improved significantly with joint incorporation of Ca, P, and Ag to the TiO₂ nanotube arrays synthesized on commercially pure Ti surface by the electrochemical method. Moreover, Eliaz and colleagues (Eliaz, 2008; Eliaz & Eliyahu, 2007; Eliaz & Sridh, 2008; Eliaz, Sridhar, Mudali, & Raj, 2005; Metoki, Liu, Beilis, Eliaz, & Mandler, 2014b; Wang, Eliaz, & Hobbs, 2011) designed a uniform, well-adhered HA-octacalcium phosphate coating on Ti-6Al-4V substrate using a one-step electrodeposition method at near-physiological conditions (37°C, pH=7.4). Meanwhile, Liang et al. (2015) applied femtosecond lasers to deposit the Ca/P phase on the Ti surface to attain better bone forming ability. Nguyen, Deporter, Pilliar, Valiquette, and Yakubovich (2004) reported that Ca/P-coated Ti-6Al-4V implants obtained by sol-gel technique displayed a trend toward a significant increase in the area of bone ingrowths across the tibiae of 17 rabbits. Recently, the Ti-6Al-7Nb titanium alloy was intensively studied for application as a medical implant because niobium improves the biocompatibility of the material in comparison with vanadium achieving the same mechanical properties (Huang, Wu, Sun, Yang, & Lee, 2014; Moskalewicz, Seuss, & Boccaccini, 2013; Oliveira et al., 2015; Rafieerad et al., 2015b). In particular, Avelar-Batista Wilson, Banfield, Housden, Olivero, and Chapon (2014) found that triode plasma nitriding of the Ti-6Al-7Nb surface followed by PVD TiN coating significantly improved the tribological properties of the alloy. Krząkała et al. (2013) demonstrated the surface modification of Ti-6Al-7Nb by PEO followed by thermal and alkali treatments of the gel-like coating formation with higher corrosion resistance in Ringer’s solution and significantly better bioactivity in simulated body fluid (SBF). Recently, some promising results for increasing biocompatibility with the bone for Ti-based implants was demonstrated by other researchers who used different methods, such as plasma-sprayed deposition, hot isostatic pressing, thermal spray, dip coating, electrophoretic deposition, etc. (Bolzonia, Weissgaerber, Kieback, Ruiz-Navas, & Gordo, 2013; Chenghao, Li’nan, Chuanjun, & Naibao, 2015; Kumar, Narayanan, Raman, & Seshadri, 2010; Li, de Groot, & Kokubo, 1996; Li, Lee, Kobayashi, & Aoki, 1996; Lo et al., 2000; Ma, Liang, Kong, & Wang, 2003; Rafieerad et al., 2015a; Vandrovcova et al., 2015; Wie, Herø, & Solheim, 1998).

As a result of our studies, calcium phosphate coatings including HA can be formed on the surface of commercially pure titanium (VT1-0) by PEO (Gnedenkov et al., 2011a; Gnedenkov, Sinebryukhov, Khrisanphova, & Scorobogatova, 2005b). Various calcium-containing salts were chosen as electrolyte components to form PEO coatings on VT1-0 titanium with a Ca/P ratio close to that of the human bone tissue (1.67). Table 1 presents the electrolyte systems, synthesis conditions, phase and elemental compositions, and Ca/P ratios for the coatings formed on titanium in electrolytes containing calcium citrate and acetate with addition of disubstituted sodium phosphate. As can be seen from the results, the coatings formed in acetate-containing electrolyte comprise HA. The coatings formed in the citrate-containing electrolyte include only TiO₂ in their composition, but judging from the value of the Ca/P ratio, calcium phosphate compounds are also present as X-ray amorphous phase. The important factor determining HA synthesis in the surface layer is the character of polarization during PEO. According to experiments, Ca₁₀(PO₄)₆(OH)₂ was obtained only in bipolar PEO mode in acetate-containing electrolyte (Table 1, sample 5), whereas in the citrate-containing electrolyte, despite high Ca and P concentrations observed in coatings (Table 1, sample 4) at low electrolyte concentrations, the calcium phosphate compounds were X-ray amorphous phases. It is likely that in the bipolar mode of treatment, repolarization of the titanium electrode leads to successive saturation of the near-electrode space with calcium and phosphate ions. (During cathode polarization, the surface electrolyte layers are saturated with Ca²⁺ ions, and during anode polarization, they are saturated with HPO42- ions, which, during interaction, form molecules of calcium-phosphate compounds Ca₁₀(PO₄)₆(OH)₂, Ca₃(PO₄)₂, and CaHPO₄·2H₂O). In addition, in the bipolar mode of PEO, a large number of electrolyte components become involved in the plasma composition, thereby providing a more intensive electrochemical synthesis of calcium phosphate compounds from the corresponding elements than in the monopolar mode. The bipolar PEO mode of coating formation has important advantages over the monopolar mode. In the interval of cathode polarization, the double (Helmholtz or Gouy-Chapman) electric layer is re-formed. During anodic polarization, this layer reduces the ion transport and the probability of penetration of oxidizing reagents into the electrode under treatment. In addition, along with disordering of the double electric layer of the electrolyte during the cathode period, the re-formation of the double layer (the space charge region) in the material of the coating takes place as well, which decreases the resistance to charge transfer. Because of these reasons, more powerful plasma discharges are observed on the anode with the subsequent positive bias. In this case, a larger number of electrolyte elements are involved in the plasma. The heat released during thermal destruction of citrate- and acetate-containing anion complexes provides a longer duration of thermal influence on the material of the oxide layer being formed (in comparison with the duration provided during destruction of light anions), which in turn facilitates the transition of the metastable state of the compounds entering the coating material to the stable one. It affects the morphology of the surface layers and changes the anticorrosive characteristics of the coatings.

Investigation of the coatings’ elemental composition demonstrates that calcium and phosphorus are present in all samples (Table 1), whereas X-ray diffraction (XRD) analysis did not confirm the presence of calcium phosphate compounds in the coatings. This means that either the concentration of such compounds is smaller than 10% (the detection limit of the XRD method) or they are included as X-ray amorphous phases. Energy dispersive X-ray spectroscopy (EDX) data allow the calcium-to-phosphorus ratio to be calculated, which demonstrates the degree of composition conformity of PEO coatings to the mineral component of the bone tissue (Gnedenkov et al., 2011a).

Morphology investigations of the coatings formed in citrate- and acetate-phosphate electrolytes through scanning electron microscopy (SEM) demonstrates that the surface is rough and porous. It seems that the pores facilitates bone tissue growth on them, leading to the formation of a stronger adhesion with the bone. The important advantage of porous coatings is the prospect of obtaining composite heterostructures. The pores and roughness of the coating surface can be filled with various medicines, e.g. antibiotics, to further decrease the possibility of inflammatory processes during implant applications (Gnedenkov et al., 2011a). The rough porous surface of the formed coating and its mineral content, along with the mineral composition and morphology of the bone tissue, suggest fast fixation of the implant because bone tissue grows through the pores on its surface.

One of the most important conditions for bone substitutes, in particular, those used as loaded ones, is the affinity of their mechanical characteristics to the properties of natural bone tissue. Titanium products have an elasticity modulus, E_Y, of 80 GPa (with microhardness, H, of 2.7 GPa), which is much greater than that of natural bone (E_Y=20 GPa; H=1.2 GPa). The discrepancy between these parameters is a reason for the occurrence of edge stresses in the junction of the implant with the surrounding bone. At certain critical strains arising under load, the bone can detach from the implants, leading to necrosis. Therefore, in the present work, we investigated the mechanical and elastic-plastic characteristics of calcium phosphate layers on the titanium surface. According to measurements (Gnedenkov et al., 2011a), the microhardness of the surface PEO layer is 2.2 GPa, which is lower than that of commercially pure titanium VT1-0 (H=2.7 GPa). At the same time, the Young’s modulus of the coating (E_Y=30 GPa) is almost 2.5-fold smaller than that of the substrate (E_Y=80 GPa). Thus, the calcium phosphate coating formed on the titanium surface has mechanical properties similar to those of natural bone tissue in comparison with the substrate, which is also an important advantage of the heterogeneous oxide PEO layers. Here, the calcium phosphate coating can function as a damper, minimizing the probability of bone detachment from the implant under strain (Gnedenkov et al., 2011a; Gnedenkov et al., 2005b; Li et al., 2007; Zheng et al., 2008).

The main factor in the successful application of any material in medicine is its preliminary testing invitro and invivo. As a rule, tests in vivo are long term, up to several months. Therefore, preliminary testing in vitro, which allows the materials to be classified according to the degree of their bioactivity under available laboratory conditions in simple experiments, is usually performed. To estimate the biological activity of samples with PEO coatings under in vitro conditions, the examined samples are immersed in SBF (Kokubo & Takadama, 2006) at a constant temperature (37°C±0.5°C). After soaking the coatings in the SBF solution saturated with Ca²⁺ and PO43- ions, mineralization of crystals of calcium phosphate compounds on surface active centers was observed together with the formation of flaky HA structure (Figure 1), whose presence was confirmed by XRD (Gnedenkov et al., 2011a; Gnedenkov, Khrisanphova, Ignat’eva, Sinebryukhov, & Zavidnaya, 2005a). The results obtained by Gnedenkov et al. (2005b) are in agreement with Huang, Xu, and Han (2005), Ryu, Song, and Hong (2008), Wei, Zhow, Jia, and Wang (2007), and Wei, Zhow, Jia, and Wang (2008), where it was stated that calcium phosphate coatings with definite roughness could induce the formation of crystalline HA. Huang et al. (2005) demonstrated that when the metal with coating was immersed in the SBF solution, HA grew the coating surface. Its formation in the SBF solution is the criterion of film bioactivity. The HA-containing coating improves the bioactivity of oxide film and stimulates biological reactions in vivo, thus accelerating bone growth and reducing the fixation time of the metal implant. Meanwhile, it is well known that PEO coatings formed on the titanium surface usually contain TiO₂ (in rutile and/or anatase crystallographic modifications) in the calcium phosphate layers (Arifin et al., 2014). The isoelectric point of rutile is 4.6±0.4; for anatase, it is 5.9±0.2 (Hanawa et al., 1998). Rutile and anatase in the SBF solution can be charged negatively and electrostatically adsorb Ca²⁺ ions on the coating surface, with the latter attracting OH^- and HPO42- ions creating supersaturation by calcium phosphate in the subsurface coating layer. Under certain conditions, calcium phosphate can be mineralized from the solution with the formation of HA (Huang et al., 2005; Ryu et al., 2008; Wei et al., 2007, 2008).

Figure 1:

SEM images of the commercially pure titanium VT1-0 surface with coating formed in acetate-phosphate electrolyte before (A) and after (B) immersion in the SBF solution for 30 days.

The efficiency of osteointegration, which depends on many factors such as bone state, composition and properties of the implanted product, adhesion of the implanted product, and the mechanical properties of the processes that take place at the interface between the coating/bone tissue and the coating/implant (Kim & Kawashita, 2003), can be only estimated in a direct experiment with a biological object. To study the bioactivity of the PEO layers in more detail, invivo experiments (Gnedenkov et al., 2011a) were carried out on the implantation of titanium implants with coatings synthesized in citrate-phosphate and acetate-phosphate electrolytes directly in laboratory mice. The results demonstrate that subcutaneous implantation of discs during 45 days produced no side effects of inflammatory and allergic origin. It was established that phosphorus and calcium compounds at definite ratio had to be present in the coating to obtain its osteoinductive properties. However, as seen from experimental results, this factor was not a unique significant parameter. Another important condition was the roughness of the calcium-phosphate layer and the absolute Ca and P concentrations in the coating (Gnedenkov et al., 2011a). The coarsely fibered bone tissue with cavities filled with bone marrow (Figure 2) was revealed in histological preparations of tissue plates grown on artificial surfaces with PEO layers painted by hematoxylin-eosin.

Figure 2:

Morphological structure of the tissue grown on calcium phosphate coatings in the test of ectopic bone formation in mice at 400× magnification.

(1) Fragments of calcium phosphate coating. (2) Bone plate. (3) Lacunas filled with red bone marrow.

The quantitative parameters of histological composition (bone and bone marrow) of tissue plates together with the roughness R_a, phase composition and Ca/P ratio of coatings are presented in Table 1. Data demonstrate that the biologically active calcium phosphate coatings are formed on Ti surface by PEO in electrolytes containing sodium citrate and acetate. In this case, XRD results, which shows that only TiO₂ was found in the composition of some coatings, do not deny the presence of calcium phosphate compounds as the X-ray amorphous phases. This can be judged from the high Ca/P ratio of samples containing only TiO₂. HA and calcium phosphate were found in sample 5 (Table 1). The Ca/P ratio was 1.92. The key factor for the estimation of the bioactive properties of the formed coatings is the amount of bone tissue formed at their surface. Investigations on the influence of the roughness parameter R_a on the quantitative characteristics invivo demonstrated that the bone area (S of bones, mm²) and bone marrow area (S of bone marrow, mm²) have no linear dependence on the roughness R_a of the artificial surfaces. However, points of extreme bioactivity were found for coating roughness in the range of 2–3 μm. Roughness is an important but is not a unique criterion for the intensive formation of bone tissue. An important role is played by both the absolute phosphorus and calcium contents in the coating and their ratio (Gnedenkov et al., 2011a). According to data (Table 1), bone formation was maximal for samples 4 and 5. Considering their different phase compositions (TiO₂, X-ray amorphous phase for sample 4, HA, and Ca₃(PO₄)₂ for sample 5), the Ca/P ratio was high in comparison with other samples presented in Table 1. Sample 3 has a sufficiently high Ca/P ratio, but the contents of Ca and P are smaller in comparison with those of samples 4 and 5. In addition, the roughness parameter R_a for sample 3 is maximal (9.13 μm) among the examined surfaces. Meanwhile, samples 1 and 2 have optimal roughness (R_a) of 2.15 and 2.86 μm, respectively (Table 1). However, the Ca and P concentrations and Ca/P ratios of these samples are much smaller in comparison with samples 4 and 5. For these reasons, it seems most likely that the growth of the bone tissue on these surface layers (samples 1–3) is low.

Thus, surface roughness is an important but not sufficient condition for bone formation (Gnedenkov et al., 2011a). Note that similar suggestions were mentioned by other authors (Aparicio, Padrys, & Gil, 2011; Aparicio, Rodriguez, & Gil, 2011; Lee, Ryu, Lee, & Hon, 2005). This is in agreement with the data of earlier invitro research that revealed the optimal roughness (R_a=2–3 μm) of artificial surfaces for manifestation of the best osteogenic properties of human multipotent mesenchymal stromal cells (MMSCs). It is possible to assume that optimal R_a values for osteogenic differentiation of human MMSC invitro and mouse MMSC in the test of ectopic bone formation are similar in many respects.

Thus, PEO has been demonstrated to allows calcium phosphate and oxide coating to be formed on commercially pure titanium VT1-0. Through the investigation of phase and elemental compositions and surface morphology, the Ca/P ratio was determined. The mechanical properties of calcium phosphate coating were established to be close to those of the bone tissue. Tests of coating in the SBF solution demonstrate the biological activity of the examined surface layers. The biological activity of the surface layers under in vivo conditions was shown to be dependent on a superposition of its parameters: chemical composition (Ca and P concentrations as well as their ratio) and morphological features (roughness).

3 Bioinert protective composite coatings

The biocompatibility of NiTi and its unusual mechanical properties make it a superior alloy for many bone implant applications, e.g. maxillofacial and dental implants, joint replacements, cervical and lumbar vertebral replacements, bone fracture anchorage and repair (Bansiddhi, Sargeant, Stupp, & Dunand, 2008; Elahinia, Hashemi, Tabesh, & Bhaduri, 2012; Fadlallah, El-Bagoury, El-Rab, Ahmed, & El-Ousamii, 2014; Razavi, Fathi, Savabi, Vashaee, & Tayebi, 2015b; Toker & Canadinc, 2014). To reduce Ni²⁺ ion release, which may cause toxic and allergic effects, an intensive investigation of titanium nickelide surface coatings was performed. Yu et al. (2015) showed that the chemical bond attachment of bioorganic poly(ethylene glycol) film on NiTi alloy surface via γ-irradiation was efficient in obtaining a bioactive material for medical applications. Chembath, Balaraju, and Sujata (2015) developed a nanogrid structure with a combination of one-dimensional channel and two-dimensional network-like patterns on the nitinol surface through application of chemical treatment. Furthermore, other promising techniques for NiTi surface modification, e.g. the electrophoretic deposition processes, hybrid cathodic arc/glow discharge plasma-assisted chemical vapor deposition, low-temperature plasma nitriding process, magnetoelectropolishing, electrochemical anodization, electrohydrodynamic spraying, sol-gel, were reported (Bakhshi et al., 2011; Czarnowska et al., 2015; Dudek, Plawecki, Dulski, & Kubacki, 2015; Gill et al., 2015; Jamesh et al., 2015; Perez et al., 2007; Saugo, Flamini, Brugnoni, & Saidman, 2015; Yang et al., 2010).

On the one hand, only a few papers in which PEO was used for coating formation on NiTi alloy have been published. In particular, Xu and colleagues (Xu, Liu, Wang, Yu, & Zhao, 2008; Xu, Liu, Wang, & Zhao, 2008; Xu, Liu, Yu, & Zhao, 2009) reported that the PEO treatment (pulsed bipolar mode) in a solution of sodium aluminate and sodium hypophosphite of biomedical nitinol results in rough and porous γ-Al₂O₃ coatings containing Ca and P with a good corrosion performance. Further (Liu, Xu, Wang, Zhao, & Shimizu, 2010), Al₂O₃-coated NiTi alloy soaked in SBF for 14 days was found to have apatite-forming ability. On the other hand, Siu and Man (2013) showed, using an AC power source, that PEO in aqueous solutions of Na₂SO₄ and NaOH provided the formation of a thick, rough, and porous TiO₂ (anatase) coating on the nitinol substrate, which facilitates HA formation. Wang, Eliaz, and Hobbs (2012) suggested the effective way to improve the bioactivity of NiTi via the TiO₂ coating using PEO in the concentrated H₂SO₄ electrolyte (DC power supply).

As a result of our search devoted to the formation of protective surface layers, the electrolyte containing sodium aluminate (NaAlO₂), sodium carbonate (Na₂CO₃), and three-substituted sodium phosphate (Na₃PO₄·12H₂O) was chosen. Based on the principles of targeted PEO synthesis designed by Gnedenkov and colleagues (Gnedenkov et al., 2007; Gnedenkov et al., 2011b; Khrisanphova, Volkova, Gnedenkov, Kaydalova, & Gordienko, 1995), it was assumed that surface layers containing aluminum compounds (oxides and phosphates) and TiO₂ could be formed in the suggested electrolyte on the titanium nickelide surface. In addition, dimethylglyoxime (DMGO), which is known to bind Ni²⁺ ions into chelate compounds (nickel dimethylglyoximate Ni(HC₄H₆N₂O₂)₂ having red color) was added to the electrolyte. It was assumed that Ni²⁺ formed in the electrolyte solution in the process of oxidation during anode dissolution of nickel from NiTi would be bound by DMGO into complex compounds stable at room temperature and insoluble in aqueous electrolyte. According to the suggested mechanism, the deposition of this complex will reduce the porosity of the newly formed anode layer, thus increasing its protective properties. The presence of nickel dimethylglyoximate in the electrolyte solution is proven by red deposits formed in the oxidation process both in the electrolytic cell and on the surface of the sample being oxidized.

According to XRD data (Gnedenkov & Sinebryukhov, 2009; Sinebryukhov, Gnedenkov, Khrisanphova, & Gnedenkov, 2009), aluminum phosphate (AlPO₄) and nickel aluminum oxide (NiAl₂O₄) are present in the coating. Analysis of XRD patterns of the surface layers of some samples shows the presence of oxygen-containing nickel and titanium compounds (Ni₃Ti₃O), whose concentrations are small. At the same time, no titanium oxides were revealed in the surface layers. Figure 3 shows an SEM photograph of the cross section of the sample with PEO coating on the NiTi surface. The thickness of the PEO coating is equal to about 30 μm. Analysis of the photograph demonstrates that the surface of the layers formed on NiTi is heterogeneous: pores and roughness can be seen. It is likely that the coatings have cluster morphology.

Figure 3:

Cross-section SEM photograph of the PEO-coating formed on the NiTi surface.

As the behavior of the implant material in a corrosion-active medium significantly depends on its surface (Gnedenkov et al., 2000; Rondelli, 1996), the titanium nickelide samples were treated by several methods:

• Deposition on the surface of superdispersed polytetrafluoroethylene (SPTFE Forum^® Institute of Chemistry FEB RAS, Vladivostok, Russia) followed by heat treatment at 100°C for 1 h.

• PEO in the potentiodynamic monopolar mode up to 180 V at a voltage sweep rate of 0.25 V/s and in the sign-alternating bipolar mode. The bipolar mode represented a combination of the anode potentiostatic (dU/dt=0.25 V/s) and cathode galvanostatic (j_c=0.5 A/cm²) polarization modes. The anode and cathode polarization periods had a τ_a/τ_c ratio of 4.

• PEO (in the mode corresponding to item 2) with subsequent deposition of SPTFE on the coating surface and heat treatment at 100°C for 1 h.

Analysis of the electrochemical behavior of samples in the physiological solution (Figure 4) demonstrated that the maximum protective properties characterized composite coatings with the PEO layer treated by SPTFE (curves 6 and 7 in Figure 4). Such treatment led to the ennoblement of the free corrosion potential (E_c) of the sample and to the decrease of the corrosion current density (j_c). Incorporation of DMGO into the forming solution also yielded a small increase in the protective properties (this conclusion was drawn based on a comparison of samples 3 and 5). Table 2 presents the comparative corrosion characteristics of the NiTi samples treated by different methods. The value of the polarization resistance (R_p) of the sample evaluated from the linear region of the polarization curve in the field of free corrosion potential is also presented in Table 1. The corrosion current of the PEO coating formed in the bipolar mode (sample 4) is reduced by 2 orders of magnitude as compared to the uncoated NiTi (sample 1). In addition, it is better than the previously reported ones for other techniques. In particular, Shi, Cheng, and Man (2007) showed that anodic oxidation of NiTi in acetic acid resulted in oxide layer formation, which provided j_c=0.2 μA/cm². Applying the laser treatment, Wong, Cheng, and Man (2007) found that thin (few dozen nanometers) thermal oxide film on NiTi was characterized by j_c=0.13 μA/cm². Vojtěch, Fojt, Joska, and Novák (2010) noted that chemical etching produced good corrosion performance (j_c=3.5 μA/cm²) in terms of nickel release. In general, the corrosion behavior depends on the thickness of the protective oxide layer. However, oxide structure (amorphous or crystalline), adherence to substrate, presence of defects, and oxide chemical composition, mainly the Ni content, play important roles as well. In addition, the obtained data are also in good agreement with those of Liu, Xu, Yu, Wang, and Zhao (2009), Liu, Xu, Yu, Wang, and Zhao (2010), Xu, Liu, Wang, Yu, and Zhao (2009), Xu et al. (2008a), and Wang, Liu, Zhang, and Wang (2012), which were devoted to PEO of titanium nickelide in aluminate solution using pulsed bipolar power supply. However, the reported works focus on making the Al₂O₃ layer on the titanium nickelide surface, whereas TiO₂ is far more biocompatible as a biomaterial.

Figure 4:

Polarization curves in the physiological solution recorded with a sweep rate, v, of 10 mV/min.

Sample notation is given in Table 2.

Analysis of the results of electrochemical investigations (Gnedenkov & Sinebryukhov, 2009; Gnedenkov, Khrisanfova, Sinebryukhov, Puz’ & Gnedenkov, 2008a) demonstrates that the surface treatment of NiTi by SPTFE insignificantly influenced the state of the electrode/electrolyte interface, just slightly reducing the anode current (Figure 4). The weak effect of such treatment is most likely due to the weak adhesion of the polymer to the metal substrate and the insufficient continuity of the formed protective layer. Treatment of the sample by PEO increased its stability in the examined medium (Gnedenkov et al., 2008b). Addition of DMGO into the electrolyte increased the protective properties of coatings (curve 5 in Figure 4). However, this difference is very small. The significant increase of just the free corrosion potential can be seen on the polarization curves (curve 5 in Figure 4). The anode current for such coating, commensurable in the initial section with that of the sample with the PEO layer formed in the bipolar mode, after 0.7 V, becomes even greater than for other samples with PEO coatings. This suggests that the chelate nickel dimethylglyoximate compound deposited into the pores of the oxide layer yields a certain increase in protective properties of the surface layer. However, its concentration in the film is not high (<10%) because no reflections corresponding to this phase were fixed on the diffraction pattern. Being a thermally unstable compound (Gallyamov et al., 2004; Hanawa et al., 1998; Li, Zhang, Li, & Qian, 2006), nickel dimethylglyoximate partially decomposes during PEO, which is known to be accompanied with high temperatures realized in short-living plasma channels as well as during thermolysis in coating regions adjacent to the plasma channel. Thus, a low nickel dimethylglyoximate content in the coating pores does not allow significant corrosion protection to be provided at large anodic biases.

It should be taken into account that the PEO coating represents the structure consisting of the outer (porous) layer and inner (nonporous) sublayer. The porous intermediate layer has, as a rule, crater-like cavities with diameter up to several micrometers. Thus, the pore sizes are somewhat greater than the size of the SPTFE powder particles used for film surface treatment. SPTFE with particle size of 0.35 μm was obtained from products of polytetrafluoroethylene pyrolysis through the gas dynamic thermal destruction method (Bouznik, Kirik, Solovyov, & Tsvetnikov, 2004; Ignatieva, Tsvetnikov, Gorbenko, Kaidalova, & Buznik, 2004). The formation of a porous surface structure can be an additional advantage of PEO because the bone tissue grows better on rough implant surface, thereby filling the pores with bioinert or bioactive composites. After the treatment of the sample with PEO coating by SPTFE, the impedance modulus significantly increased and was 10 times greater than that for the sample without coating (Sinebryukhov et al., 2009). This demonstrates that SPTFE treatment allows the coating pores to be filled with polymer, with the additional barrier layer preventing metal ion penetration into the solution to be created. This conclusion was confirmed by the atomic absorption method (Puz’, 2006). The concentration of nickel ions in the SBF solution as a result of immersion in this media for 2 weeks (37°C) of PEO- and SPTFE-pretreated NiTi samples has abated by 1 order of magnitude in comparison with bare (without treatment) titanium nickelide (Table 3). Note that this is in a good agreement with the results reported by Wang et al. (2012) for the nickel release rate in Hank’s solution at 37°C.

Thus, the prospects of the application of PEO have been demonstrated for the formation of coatings on the titanium nickelide surface, improving its morphological structure and electrochemical properties. Application of SPTFE as part of the composite coatings on the given material increased its stability in a corrosion-active medium (Gnedenkov et al., 2011b). The combined impact of polarization and plasma on the sample surface during PEO allows formation of coatings with protective properties. Note that after filling pores with SPTFE powder, thermal treatment is required to obtain smoothed hydrophobic surface layers (Gnedenkov, Sinebryukhov, & Sergienko, 2006; Gnedenkov et al., 2011b). In this case, the polymer fills pores, into which the medicine can be preliminarily injected, and thus, their diffusion from the surface layers can be reduced. This increased the duration of the therapeutic effect of the medicine deposited in the pore. In addition, the bioinert coating being formed considerably reduced diffusion of nickel ions from NiTi and hindered their accumulation in human tissues (Gnedenkov & Sinebryukhov, 2009; Gnedenkov et al., 2011b).

4 Corrosion-resistant, bioactive composite coatings

Mg alloys are suitable for medicine applications such as orthopedic implants and stents for vessel dilatation due to their biodegradability (fully dissolve in the human body), which prevents secondary surgery for implant removal (Narayanan et al., 2015a,b). However, magnesium corrodes quickly in physiological environments, limiting the complete recovery of organism. Thus, the anticorrosion coating on Mg implant surface is required to control the rate of their degradation. Recently, Razavi et al. (2014, 2015b) and Razavi, Fathi, Savabi, Vashaee, & Tayebi (2005a) showed that modification of the surface of AZ91 using bredigite (Ca₇MgSi₄O₁₆), diopside (CaMgSi₂O₆), and fluoridated HA (Ca₁₀(PO₄)₆OH_2-xF_x) coating through the combination of anodic spark deposition or PEO and electrophoretic deposition techniques resulted in a significant enhancement of the magnesium implant biocompatibility. At the same time, Santos et al. (2015) demonstrated, using a parallel nano-assembly route, the one-step fabrication of a novel graphene oxide/HA nanoparticles/phosphate coating for better biocompatibility and osteointegration of ultrahigh-purity Mg implant. Meanwhile, Asi, Nemeth, and Tan (2014) reported the approach in the formation of protective and bioactive coating on AZ31 magnesium substrate due to the hydrothermal monetite (CaHPO₄) film deposition. Furthermore, at present, Mg-Ca, Mg-Zn-Ca, AM60, Mg-Gd-Ca-Zr, Mg-Ca-Sr, Mg-Li-Al-Ce, etc. had been widely investigated as biodegradable medical implants (Atrens, Liu, & Abidin, 2011; Bakhsheshi-Rad et al., 2014; Bornapour, Celikin, & Pekguleryuz, 2015; Chen, Xu, Smith, & Sankar, 2014; Leeflang, Dzwonczyk, Zhou, & Duszczyk, 2011; Pan et al., 2015; Shi et al., 2015). Despite intensive research and designing of a number of promising techniques, the corrosion resistance of coated biodegradable Mg alloys is still insufficient for their practical application in medicine.

For our investigations, rectangular specimens (5×30×1 mm) made from wrought MA8 magnesium alloy (1.5–2.5 wt% Mn; 0.15–0.35 wt% Ce; Mg – balance) were used as substrates for PEO coating formation. The specimens were preliminarily polished to achieve a surface roughness R_a of 0.12 μm, washed with distilled water, and degreased with ethanol. The electrolyte was prepared via successive dissolution of 25 g/l (C₃H₇O₆P)Ca·2H₂O and 5 g/l NaF in distilled water under constant stirring, with pH adjusted to 10.9–11.3 by adding 20% NaOH solution. Post-treatment of the PEO coating using SPTFE was suggested to decrease surface roughness, thus allowing further reduction of the severity of corrosion attack. SPTFE powder was applied to the PEO-coated surface by triboelectric method, which consists of mechanical rubbing of SPTFE on the PEO surface.

The PEO coating fabricated on the magnesium alloy in the calcium glycerophosphate-sodium fluoride electrolyte was found to consist of crystalline magnesium oxide and HA phases as well as an amorphous constituent. According to cross-sectional analysis, the coating is approximately 60 μm thick. The coating surface was found to have a highly convoluted heterogeneous morphology (Figure 5). Pores and cracks typically emerge due to high pressure and temperature realized in the plasma channels during the PEO process (Gnedenkov et al., 2010).

Figure 5:

SEM images of the PEO-coated Mg-Mn-Ce magnesium alloy.

EDX analysis confirmed that the coatings contained such elements as Ca, P, Mg, Na, O, C, and F. The Ca/P ratio in the coating is equal to 1.61, which is close to that in the bone tissue (1.67). According to Sun, Lu, Yuan, Jing, and Zhang (2011), the formation of a binary compound consisting of sodium and magnesium fluorides NaMgF₃ in an amorphous form is also possible in the coating.

The presence of carbon within the coating can be attributed to the organic components of the electrolyte (namely, calcium glycerophosphate) that were adsorbed at the surface and/or failed to combust in discharge channels during the PEO process. Organic radicals facilitate a longer-lasting glow in the plasma discharge, which results in high-temperature synthesis of crystalline compounds (in this case, HA) within the coating. For this reason, an increased Ca₁₀(PO₄)₆(OH)₂ content was observed in the regions where carbon content was lower compared to other regions. The maxima of the calcium and phosphorus concentrations in accordance with EDX data obtained on the cross section were found in the outer part of coating, whereas the inner part attached to substrate consisted of Mg, O, and F. The composition of this part of coatings formed in the fluoride-containing electrolytes consists predominantly of MgO and low amorphous MgF₂ (Sun et al., 2011). According to the results of both in vitro and in vivo experiments (Ma et al., 2014; Qiu, Wan, Tan, Fan, & Yang, 2014; Zhang, Dai, Wei, & Wen, 2012), HA-containing PEO coatings fabricated on magnesium alloys exhibit high biological activity, considerably accelerating the growth of the bone tissue at the implant surface.

According to scratch test results, the adhesion of these coating to the substrate is very high. The critical load for such coating is equal to 22.8±1.5 N, which met the requirements for medical implant coatings. Moreover, this coating, according to impact test data, does not peeled with falling weight (1 kg) from a height of 1 m (Gnedenkov et al., 2013, 2014b). The impact test was carried out according to ISO 6272 using a ZIT 2440 Impact tester (Zehntner GmbH Testing Instruments, Sissach, Switzerland).

The parameters of corrosion resistance derived from potentiodynamic polarization curves and Bode diagrams (impedance modulus |Z| and phase angle θ as functions of frequency f) are listed in Table 4. Analysis of the data in Table 3 affirms that PEO coatings considerably reduce the dissolution rate of magnesium in the active region. Some quantitative differences in |Z|_{f→0 Hz} and R_p values shown in Table 4 link to the polarization conditions of the sample under investigation used at potentiodynamic polarization and impedance measurements. The impedance modulus at low frequencies, which characterizes the coating protective properties, is |Z|_{f→0 Hz}=5.07×10⁴ Ω·cm². This value is over 60-fold higher than that for the uncoated sample (|Z|_{f→0 Hz}=8.09×10² Ω·cm²). The results are in a good agreement with a previous report (Razavi et al., 2014).

The results of corrosion evaluation in Hank’s balanced salt solution (at 37°C) conducted using potentiodynamic polarization and EIS techniques are presented in Table 5. As the magnesium alloy substrate has higher stability in the Hank’s solution (which has lower concentration of chloride ions) even at 37°C than the 3% NaCl solution at 20°C (Tables 4 and 5), the corrosion resistance of the PEO-coated alloy is not significantly affected by the physiological medium (Table 5) (Gnedenkov et al., 2013, 2014b).

To further improve the anticorrosion protection of the alloy, SPTFE post-treatment of the PEO coating was applied, followed by thermal treatment. SPTFE (unlike polytetrafluoroethylene) is a convenient material for the formation of composite coatings due to its properties. This material contains low-molecular fractions, which have softening and melting temperatures lower than those of polytetrafluoroethylene. According to thermal analysis data, the melting point of SPTFE is about 220°C, which leads to the penetration of SPTFE into the pores of the PEO layer. This particular post-treatment procedure ensures sealing of predominant pores and cracks (Figure 6A and B) (Gnedenkov et al., 2014a,b; Minaev et al., 2011). The comparison of surface SEM images of the PEO coatings before (Figure 5) and after SPTFE treatment (Figure 6A) indicates smoothing of the surface morphology as a result of the SPTFE post-treatment.

Figure 6:

SEM images of the Mg-Mn-Ce alloy with PEO coating after filling of pores with SPTFE.

(A) Surface observations. (B) Cross-section observations.

From EDX analysis, fluorine is predominantly distributed in the porous part of the coating (Figure 6B). Thus, the fluorine content in points 2 and 3 is equal to 23 and 36 at%, respectively, which is considerably higher than that in the point 1 (about 8 at%). According to data in Table 5, such post-treatment significantly reduces the corrosion current density (>3 orders of magnitude), substantially (by about 3 orders of magnitude) increasing both polarization resistance and impedance modulus. The graphs of the dependence of the impedance modulus on frequency corroborate the conclusions made based on the analysis of polarization curves. The values of the impedance modulus measured at low frequency (|Z|_{f→0 Hz}, Table 5) for the samples with a polymer layer on the surface are several orders of magnitude higher than those of the PEO coating not treated with polymer or the bare substrate itself. The results indicate that PEO with SPTFE coating of Mg alloys manifests higher values of R_p and |Z|_{f→0 Hz} than other surface treatment methods such as duplex PEO/molybdate conversion coating of Mg-Li alloy (Sun e al., 2011), silicon-doped calcium phosphate coating of AZ31 alloy via pulse electrodeposition (Qiu et al., 2014), dicalcium phosphate dihydrate coating of magnesium alloy by PEO coupled with hydrothermal treatment (Chang, Tian, Liu & Duan, 2013), and others (Niu et al., 2013; Shi, Xu, & Wang, 2010; Tan et al., 2010).

Thus, the HA-based PEO coatings with subsequent SPTFE treatment substantially reduce the corrosion rate of the magnesium alloy substrate. The original method for surface treatment of Mg has a good potential for application in biodegradable magnesium implants, which could bring the implant surgery to a qualitatively new level.

5 Conclusions

PEO is an efficient, low-cost, environmentally friendly, and attractive surface modification technology. It is a unique method of forming functional coatings for a number of intensively growing, specific state-of-art applications, e.g. medicine. At present, PEO is in a transition stage from research to commercialization, and it is necessary to further study the fundamentals of this technique to advance scientific understanding and design new functional coatings for high-technology fields. This review presents the advances associated with functional coatings formed by PEO for enhancement of the performances of medical implants (made from commercially pure titanium, titanium nickelide, and Mg-Mn-Ce magnesium alloy). It was shown that formation of calcium phosphate surface layers as a result of surface modification of the commercially pure titanium increased its biological activity, which was confirmed viain vitro and in vivo experiments. The review is also focused on the availability the PEO for preparing bioinert polymer-containing composite coatings on the NiTi shape memory alloy to reduce significantly the release of Ni²⁺ ions into the organism. The suitability of composite coatings containing HA formed by PEO to provide bioactivity as well as to control the corrosion degradation of resorbable Mg-Mn-Ce magnesium alloy implant was demonstrated. The critical factors determining coating performances, such as surface chemistry, corrosion rate, and morphology, are identified and reviewed. Finally, perspectives for the PEO and future trends are reported.