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Coatings for metallic biomaterials
Coatings for metallic biomaterials
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Toshihiro Kasuga Division of Advanced Ceramics, Nagoya Institute of Technology, Nagoya, Japan
14.1
Introduction
The high-performance biomaterials for replacing damaged hard tissues, such as artificial joints, bone fillers, dental implants, and so on, are in high demand among the aged people. Metallic biomaterials with biological safety, such as stainless steel, cobalt (Co)-based alloys, titanium (Ti) and its alloys, are reasonable for replacing hard tissues at load-bearing sections. Ti and its alloys are especially widely used as orthopedic and dental implants because of their excellent biocompatibility. Much attention has been paid to β-type Ti alloys because they have high mechanical strength with low modulus of elasticity. One of the attractive metals is β-type Ti alloys, which contain elements sufficiently tolerated by living tissues, such as niobium (Nb), tantalum (Ta), and zirconium (Zr) (Niinomi, 2003; Kuroda et al., 1998). A Ti-29Nb-13Ta-4.6Zr (denoted by TNTZ) alloy, which has been suggested as one of the best materials designed, shows a high tensile strength of 1 GPa and low Young’s modulus of 60–80 GPa. Although Ti and its alloys do not make a chemical bond with living bone, they can adhere to it with almost no fibrous tissues. Calcium phosphate ceramics such as hydroxyapatite (Ca10(PO4)6(OH)2: HA) and β-tricalcium phosphate (β-Ca3(PO4)2: β-TCP) exhibit bioactivity, which is the ability to make a bond chemically with natural bone tissues (Hench, 1992). Some glasses and glass-ceramics are also well known as bioactive materials. Some portions in artificial substitutes for hard tissue may require bioactivity. Bioactive surface modification is, in general, applied to Ti and its alloys to further improve their bio-functionality, including osteoconductivity. The bioactive ceramic coatings prepared using various processes have been investigated so far and some materials are clinically used as biomaterials for dental and surgical implants. In this chapter, some bio-functional coatings on metals are briefly reviewed.
14.2
Calcium phosphate ceramic coatings
Implants coated with calcium phosphates such as HA and β-TCP have been widely used in orthopedics and dentistry. The cementless fixation technique combines the mechanical strength, ductility, and ease of fabrication with bioactivity. Since much earlier, calcium phosphate coatings have been applied by a variety of methods, such Metals for Biomedical Devices. https://doi.org/10.1016/B978-0-08-102666-3.00014-6 © 2019 Elsevier Ltd. All rights reserved.
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as dip-coating (Li et al., 1996), electrophoretic deposition (EPD) (Ducheyne et al., 1986), plasma spraying (Lacefield, 1993), ion-beam sputtering (Ong and Lucas, 1994), flame spraying (Berndt et al., 1990), wet-process coating, and so on. In this section, some recent possible coatings are briefly discussed.
14.2.1 Thermal sprayed coatings Thermal spraying employs a plasma/ionized gas or fuel/oxygen (O) gas to melt the ceramic particles and subsequently to carry them onto the surface of substrates. Two important properties of thermal sprayed coatings are: (i) bonding strength between the calcium phosphate layer and the metallic substrate, and (ii) dissolution behavior in body fluid. Residual stress arises at the interface between the metallic substrate and the coating layer due to the large difference in their thermal expansions. Amorphous calcium phosphate tends to dissolve rapidly in body fluid: coatings with low crystallinity rapidly become weak and may induce inflammatory responses. A relatively high bonding strength on Ti or its alloys was reported by Tsui et al. (1998); the values were in the range of 20–40 MPa after optimization of coating properties. The crystallinity of HA in their coating layers was relatively high (83%–94%). Their high crystallinity was very important for the long-term benefit of the coated implants. A radio-frequency thermal plasma spraying (rf-TPS) method was developed to obtain strong bonding between the HA coating and Ti substrate, employing an HA/ Ti composite coating (Inagaki et al., 2003). The HA/Ti composite coatings were prepared by controlling the starting composition using two microfeeders, which fed the HA and Ti powders at an accurate rate. The starting powders were fed so as to form the coated layer with a compositional gradient from a Ti-rich phase at the bottom to an HA-rich at the top. The composition of plasma gas played an important role in improving the bonding strength of the coatings. Rf-TPS with nitrogen (N)-containing plasma gas induced the formation of the coatings with excellent bonding to the Ti substrate (the tensile bonding strength of 65 MPa). The microstructure of the interface between deposited Ti particles in the coatings was significantly influenced by plasma gas composition during rf-TPS. By rf-TPS with N- or O-containing plasma gas, Ti nitride or Ti oxide was formed in the deposited Ti particles, respectively. The bonding of the composite coatings on the Ti substrate was suggested to be influenced by microstructure at the interface between Ti particles. Surgical Site Infection (SSI) is one of the complications in artificial joint replacement surgery, and its treatment is often not easy. About 50% of the infecting organisms are Methichillin-resistant Staphylococcus aureus (MRSA) and methicillinresistant Staphylococcus epidermidis (MRSE). The SSI associated with the surgery, which is called “implant related infection” or “peri-prosthetic joint infection”, is known to be linked to the formation of biofilm and it has a serious influence on the outbreak and severity of infectious disease. Antibacterial drugs don’t work well on bacteria in biofilms, and the treatment of SSI is difficult due to an increase in drug-resistant bacteria such as MRSA and MRSE. A silver (Ag)-containing calcium phosphate (CP) coating has been prepared using frame spraying to reduce the incidence of SSI in total hip arthroplasty (THA) (Noda et al., 2009). A mixture of 97% HA and 3% Ag2O powders (in wt%) was
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flame-sprayed on metallic Ti using an acetylene torch to be coated with a compositionally homogeneous amorphous calcium phosphate layer of 40 μm thickness (denoted by Ag-CP coating, hereafter). As shown in Fig. 14.1A and B, a typical surface structure of thermal spraying was obtained. Energy dispersive x-ray spectrometry (EDX) showed that the content of Ag2O in the sprayed layer was reduced to 0.3%. However, its calcium (Ca)/phosphate (P) ratio was almost unchanged from the nominal composition. The amount of Ag+ ions dissolved from the Ag-CP coating in fetal bovine serum (FBS) was 78 ppb after 1 h of soaking and 260–330 ppb after 24–48 h, respectively; Ag+ ion dissolution at an initial stage would be important for preventing SSI (Noda et al., 2009).
Fig. 14.1 Scanning electron microscopic (SEM) images of (A) surface and (B) cross-section of Ag-CP coating. No silver-related particles were observed. Reprinted from Noda, I., Miyaji, F., Ando, Y., Miyamoto, H., Shimazaki, T., Yonekura, Y., Miyazaki, M., Mawatari, M., Hotokebuchi, T., 2009. Development of novel thermal splayed antibacterial coating and evaluation of release properties of silver ions. J. Biomed. Mater. Res. B Appl. Biomater. 89B, 456–465. Copyright 2008, with permission from Wiley.
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After phosphate buffer solution (PBS) containing 102 CFU of MRSA was injected in a hole drilled into the tibial medullary cavity of 10-week-old Sprague-Dawley rats, Ag-CP coating and conventional CP flame-sprayed coating (denoted by CP coating, hereafter) samples were implanted and closed. After 24, 48, and 72 h, the tibia including the implant was dissected out and then crushed and ultrasonically irrigated with PBS. The dilution of this PBS was incubated on agar plates at 37°C and the number of inoculated viable cells was calculated (a dilution plate method). It was confirmed that the number of bacteria in the case of Ag-CP coating was significantly smaller (P < .02) than that in the case of CP coating (Akiyama et al., 2013). The silver amounts in the serum of artificial joint replacement surgery using Ag-CP coating showed a maximum value at day 14 postoperatively (average value 1.4 ng/mL), and then they gradually decreased. During the experimental period, no troubles caused by argyria were observed; high safety was indicated (Eto et al., 2016). Osteoconductivity of the Ag-CP and CP coatings was evaluated in tibia models using rats and rabbits. Newly formed bone was observed on the surface layer of both the samples, and there was no significant difference between them. Both of them were concluded to have excellent osteoconductivity. The Ag-CP coating is emphasized to be a great coating technology for the formation of functional surface with both osteoconductivity and antibacterial activity. The coating has obtained the clearance of Japanese Ministry of Health, Labor and Welfare in 2016 for femoral stem and acetabular cup of THA.
14.2.2 Thin film coatings Physical vapor deposition (PVD) is one of the most promising methods to obtain thin, uniform, and dense CP coatings on metallic substrates. Radiofrequency (rf ) magnetron sputtering has been used widely for coatings of thin films with excellent bonding to the substrates and has been applied to the coatings of CP films on commercially pure Ti (cpTi) (Yoshinari et al., 1997; Wolke et al., 2003) and Ti alloys (van Dijk et al., 1995), and the method has a great advantage in that the coatings can be achieved at relatively low temperatures. That is, no degradation in the mechanical properties of the metallic substrates occurs. Narushima et al. (2005) prepared CP films on cpTi substrates by rf magnetron sputtering using a hot-pressed, dense β-TCP target. The resulting films were dense and uniform with a smooth surface, consisting of amorphous CP and/or oxyapatite phases (Fig. 14.2; Ueda et al., 2007). The bonding strengths between the coating and the cpTi substrate were estimated to be >60 MPa. When the film was soaked in Hank’s balanced salt solution (HBSS) or PBS, the dissolved amount of Ca2+ ions from amorphous coatings was significantly larger than that from oxyapatite coatings. Animal test using beagle dogs showed that cpTi coated with the CP film was fixed rapidly and strongly with bone (Narushima et al., 2007; Ueda et al., 2007). CpTi with a sputtered HA thin layer was developed as dental implants in Japan. Animal tests using canine mandibles for 28 weeks showed excellent bone-to-implant contact (BIC) by scanning electron microscopic analysis: the values of HA-sputtered implants showed larger (98.1%), in comparison with those of machined implants
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Fig. 14.2 Cross-sectional SEM image of calcium phosphate film on a cpTi plate prepared by rf magnetron sputtering. Reprinted from Ueda, K., Narushima, T., Goto, T., Katsube, T., Nakagawa, H., Kawamura, H., Katsube, T., 2007. Evaluation of calcium phosphate coating films on titanium fabricated using RF magnetron sputtering. Mater. Trans. 48, 307–312. Copyright 2007, with permission from The Japan Institute of Metals.
(without HA-coatings) (70.4%) (Madi et al., 2016). The HA coating layer was almost completely dissolved. By evaluating the type and dimensions of bone defects after progressive peri-implantitis in dogs, peri-implant defects around the HA-sputtered implants were smaller than those around plasma-sprayed HA-coated implants (Madi et al., 2014). The dental implant system using HA-sputtered cpTi has been commercialized in Japan since 2010. Chemical vapor deposition (CVD) also has great advantages in the preparation of CP thin films coated strongly onto metallic substrates, because of high deposition rates with excellent microstructure controllability. Some reports have been published on the preparation of β-Ca2P2O7 film (Allen et al., 1996) and its conversion to β-TCP phase by postheat treatment. A few works on the direct preparation of bioactive CP films such as HA or β-TCP have been reported so far (Darr et al., 2004; Sato et al., 2007a). Bioactive CP films have been prepared on cpTi substrates by metal-organic chemical vapor deposition (MOCVD) using bis-dipivaloylmetanato-calcium (Ca(dpm)2) and tryphenyl phosphate ((C6H5O)3PO) precursors (Sato et al., 2007a). A dense HA film was prepared by controlling Ca/P atomic ratio and substrate temperature at 700–800°C. The resulting HA film showed (002) preferred orientation parallel to the substrate surface. The film can be deposited at high rates of 4–6 nm/s. HA formed in HBSS at 37°C within 6 h on the HA film. The film on cpTi was expected to show excellent bioactivity. Tsutsumi et al. (2010) prepared a polycrystalline HA film on a TNTZ surface using MOCVD by controlling the heating temperature, as shown in Fig. 14.3. The thickness of the HA film on TNTZ could be varied by 1–7 μm with the deposition time. An α-phase precipitated in the TNTZ matrix after heating the substrate during the deposition process, and the mechanical properties of HA-coated TNTZ were improved by maintaining its low Young’s modulus. MOCVD would be one of the most effective methods for HA coating on TNTZ. Calcium titanate (CaTiO3) film has also been investigated as one of the promising bioactive coatings, using a sputtering (Asami et al., 2007) or CVD (Sato et al., 2007b)
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Fig. 14.3 SEM images of surfaces of (A) TNTZ, (B) surface, and (C) cross-section of HA coated on TNTZ using MOCVD. HA deposited on TNTZ had a dense granular microstructure and its grain diameter was 2 μm. Reprinted from Tsutsumi, H., Niinomi, M., Nakai, M., Gozawa, T., Akahori, T., Saito, K., Tu, R., Goto, T., 2010, Fabrication of hydroxyapatite film on Ti-29Nb-13Ta-4.6Zr using a MOCVD technique, Mater. Trans. 51, 2277–2283. Copyright 2010, with permission from The Japan Institute of Metals.
method. HA formed on the film after soaking in HBSS for 3–7 days. It was proposed that the CaTiO3 film would firstly dissolve into the solution to locally increase Ca2+ ion concentration, where PO34 ions could be attracted and reacted with Ca2+ ions, resulting in the formation of HA.
14.2.3 Apatite coatings using wet-processing At the interface between bioactive ceramics and bone in living body, an apatite layer can be often seen: the ceramics bond to living bone through the apatite layer (Kitsugi et al., 1986), which consists of a hydroxycarbonate apatite (HCA) with small crystallites and defective structure, precipitated by incorporating carbonate anions from body fluid within amorphous CP phase. HCA has a similar composition and structure to bone. The apatite layer can form on the ceramics, metals, and polymers when they are soaked in simulated body fluid (SBF) with ion concentrations nearly equal to those of human blood plasma (Kokubo and Takadama, 2006). Two indispensable conditions needed for the HCA formation on materials using SBF are: (i) the existence of surface functional groups that induce the nucleation of apatite (e.g., Si-OH, Ti-OH, Nb-OH, Ta-OH, COOH, and PO4H2 groups), and (ii) the increase in the supersaturation concerning apatite around the surface of the materials. Metallic Ti and its alloys are generally covered with a passive layer of Ti oxide. Since the surface of the layer includes the Ti-OH group, there exists the possibility of HCA formation. However, relatively high supersaturation concerning HCA is usually needed around the surface. To increase the supersaturation in SBF, a large amount of Ca2+ ions should be supplied. For example, soluble Ca carbonate (vaterite) was expected to supply both Ca2+ and carbonate ions when soaked in SBF: the supersaturation concerning HCA increased in SBF and rapid HCA-formation occurred (Hashimoto et al., 2007). However, the mechanically adhesive strength of the HCA layer on metallic substrates might be weak.
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Fig. 14.4 Cross-sectional SEM and element distribution images of HA-coating on Mg prepared by a chemical solution deposition method at pH 8.9. The sample was scratched off from the substrate. A thin Mg(OH)2 layer formed at the bottom of the coating. The intensity of Ca and P distribution was almost constant in the outer HA-layer packed densely. Reprinted from Tomozawa, M., Hiromoto, S., 2011b. Microstructure of hydroxyapatite- and octacalcium phosphate-coatings formed on magnesium by a hydrothermal treatment at various pH values. Acta Mater. 59, 355–363. Copyright 2010, with permission from Elsevier.
Metallic Mg or its alloys have been investigated as a biodegradable material for applications in medical implants such as vascular stents. In application to orthopedic devices such as fracture fixation materials and artificial bones, there exists a serious issue with Mg corrosion, which proceeds faster than healing of the diseased area. Therefore, in order to suppress the corrosion of the Mg or its alloys, a chemical solution deposition method for developing well-crystallized CP coatings on them, has been reported (Hiromoto and Yamamoto, 2009). An HA coating was prepared by soaking in C10H12CaN2-Na2O8 (Ca-EDTA) KH2PO4 (their concentration: 0.25 mol L 1) adjusted to pH 9–12 at 90°C for 2 h. The coating showed a double-layer structure, consisting of a continuous inner layer of 1 μm thickness and an outer layer of rod-shaped HA crystals grown on the inner layer to a length of 4 μm, as shown in Fig. 14.4. Almost no nanopores were observed in the inner layer, and the coating was dense at its micrometer level (Tomozawa and Hiromoto, 2011a). From the anodic polarization curve of HA-coated or noncoated pure Mg in 3.5 mass% NaCl aq., it was shown that the cathode current density of the coated sample was much smaller than that of the noncoated one (Tomozawa and Hiromoto, 2011b). It has been shown that the corrosion of Mg or its alloys could be suppressed drastically by the dense HA coating.
14.2.4 Calcium phosphate glass-ceramic coatings Since glass-based materials can be allowed some latitude in choice of composition, they have a significant advantage in that they are controllable to improve their physico-chemical properties and microstructures after their crystallization to attain sufficient performance as biomaterials. Silicate-based glass and glass-ceramics such as Bioglass and Cerabone A-W are well-known as bioactive materials (Hench, 1992).
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In some types of glass, anionic groups can be connected through a cation, which should usually be a network modifier, to form the glassy state; such glass types are called “invert glasses”. Kasuga (2005) reported silicate-free invert glass with compositions of high CaO and low P2O5 contents; the glass consisted of PO34 (orthophosphate) and/or P2O47 (pyrophosphate) groups without PO3 (metaphosphate) group, and the phosphate groups were connected through Ca2+ ions. The proposed mother composition was 60CaO-30P2O5-3TiO2-7Na2O (in mol%, denoted by PIG hereafter). It is usually difficult to obtain glass with Ca/P > 0.75 in the binary system, but they can be prepared with small amounts of TiO2 and Na2O. TiO2 addition into phosphate glass improved their glass-forming ability and chemical durability. The structural role of TiO2 in the CP invert glass was clarified by spectroscopic analysis and molecular dynamics simulation (Maeda et al., 2017). PIG was pulverized to <10 μm in diameter (the average size; 1.0–1.5 μm) The glass powder slurry was dip-coated onto a sandblasted TNTZ and then dried. Then the substrate with the glass powder layer was heated in air at 800°C for 1 h and cooled to room temperature. Due to sintering of the glass powder, a layer with 10 μm thickness was formed on TNTZ. Many pores of several micrometers in size were seen. There existed a reaction zone of 2–4 μm in thickness between the glass-ceramic layer and the substrate. The reaction layer was spontaneously developed on TNTZ by heating in air to relax the thermal stress generated in the coating layer and the substrate, resulting in the formation of the strong joint. The coating layer in the sample consisted predominantly of β-Ca3(PO4)2 (β-TCP) phase with trace amounts of β-Ca2P2O7 (β-CPP) and TiO2 (rutile phase) crystalline phases. The tensile bonding strength of the glass-ceramic layer on TNTZ was estimated to be 26 MPa. The fracture occurs in the glass-ceramic layer with porosity; the bonding strength at the interface layer between TNTZ and the glass-ceramic was higher than the tensile strength of the glass-ceramic (Kasuga et al., 2012). The oxidized layer at the TNTZ surface was estimated to be <0.5 μm in depth. This thin oxidized layer worked to form the strong bond to PIG glass-ceramic. Aging treatment of TNTZ at 400°C is effective for increasing its hardness, modulus of elasticity, and strength (Niinomi et al., 2002). This improvement is due to the precipitation of ω-phases in the matrix β-phase during the aging. After the glassceramic-coated TNTZ was heat-treated for aging, tensile bonding strength of the sample with the coating layer of 5 μm thickness was not degraded, while that of 20 μm thickness was degraded due to the generation of microcracks, which originated from the formation of ω-phase, at the interface between the layer and the substrate (Niinomi et al., 2003; Akahori et al., 2005). The coated samples were machined for fatigue tests according to ISO1099-75 specifications (Li et al., 2004). No delaminations of the coating were observed during the fatigue test. The fatigue resistance of the samples was improved due to an increase in resistance to fatigue crack initiation caused by the formation of the α-phase during heating for the coating and the precipitation of the ω-phase from the substrate during the slow cooling. In addition, aging at 400°C for 72 h greatly increased fatigue limit of the coated TNTZ and effectively relieved the detrimental effect of sandblasting on fatigue properties.
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Fig. 14.5 Contact microradiographs around the glass-ceramic-coated TNTZ rods (5 mmϕ) implanted into femurs of Japanese rabbits after 1 month, 1 year, and 5 years. The scale bar indicates 50 μm. Reprinted from Kasuga, T., Hattori, T., Niinomi, M., 2012. Phosphate glasses and glassceramics for biomedical applications. Phosphorous Res. Bull. 26, 008–015. Copyright 2012, with permission from Japanese Association of Inorganic Phosphorus Chemistry.
A rod-shaped sample of 5 mm diameter 10 mm length was implanted into the femur of a Japanese rabbit. The sample was autoclaved in water at 121°C for 1 h. Fig. 14.5 shows contact microradiographs (CMR) after 1 month, 1 year, and 5 years of implantation. No significant change in the thickness of the glass-ceramic coating layer after the implantation could be seen. The observation showed new bone formation around the autoclaved glass-ceramic-coated TNTZ after 1 month of implantation and the bone tissue showed direct contact with implants. After 1 year and 5 years of implantation, the glass-ceramic-coating was surrounded by maturated bone tissue and had bonded directly with them. When the samples were sliced and polished for the observation after 1 year of implantation, no crack occurred between the coating and bone and also between the coating and TNTZ. This result shows excellent bioactivity of the PIG glass-ceramic coating on TNTZ (Kasuga et al., 2012).
14.3
Bioactive surface prepared by chemical treatments
Bioactive surface modifications described in this section might not be included in the category of coatings on metallic substrates. However, they are briefly reviewed, since they are very important, practical techniques for preparing bioactive surfaces on metallic substrates.
14.3.1 Chemical treatment with a concentrated alkaline solution Ti and its alloys react with alkaline aqueous solutions. The reacted surface consists of negatively charged HTiO3 nH2O, which incorporates alkaline ions. Kokubo et al. prepared amorphous sodium titanate on cpTi by soaking the metal in an aqueous NaOH l
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solution of 5 M at 60°C for 24 h and subsequent heating at 600°C for 1 h. A surface consisting of a porous layer was obtained. The sodium titanate phase was integrated with the metal substrate through a graded structure within a depth of 1.5 μm from the surface. When the treated sample was soaked in SBF, HCA formed on the porous sodium titanate layer. Since the HCA showed a graded structure, it achieved a tight adhesion to the metal substrate (tensile bonding strength; >30 MPa). When Ti alloys, such as Ti-6Al-4V, Ti-6Al-2Nb-Ta and Ti-15Mo-5Zr-3Al, were treated with NaOH and subsequent heating, they also formed sodium titanate phase, which was free of the alloying components. These treatments had no negative impact on the strength under tensile loading nor fatigue strength under cyclic loading of the metals in saline solution (Kokubo et al., 2009). The treated Ti metal and its alloys have been investigated in some animal models. When a cylindrical rod of the treated cpTi was implanted into the intramedullar canal of a rabbit femur, apatite formed on its surface within 3 weeks and it was completely covered with bone within 12 weeks. The rectangular sample implanted into a rabbit tibia formed an HCA layer on its surface and bonded to bone through the layer after 8 weeks. Surface modified Ti and its alloys started to be clinically used as artificial hip joints since 2007. Modification of a TNTZ surface using this type of treatment was also examined; HCA-forming ability was lower on the surface of alkali-treated TNTZ than that on the alkali-treated cp-Ti (Akahori et al., 2007). The HCA-forming ability might be affected due to the existence of the sodium niobate film, which was considered to delay the exchange between the Na+ and H3O+ ions. The surface modification technique to induce bioactivity can also be applied to tantalum (Ta) metal with excellent fracture toughness and malleability (Miyazaki et al., 2000). Typical treatment conditions were subjected to NaOH aq of 0.5 M at 60°C for 24 h and subsequently heated at 300°C. After the treatment, amorphous sodium tantalate phase formed on its surface. Since Na+ ions in the phase on Ta were exchanged with the H3O+ ions in SBF to form Ta-OH groups for inducing apatite nucleation, HCA formed on the treated Ta. This metal also formed apatite in the living body, and bonded tightly to bone through the apatite layer (Kato et al., 2000).
14.3.2 Surface modification by micro-arc oxidation and chemical treatment Since metallic Zr and its alloys show excellent physico-chemical properties similar to Ti and small artifact effect in MRI, they are expected to perform well as new biomaterials. It is known that Ti and its alloys show high biocompatibility close to bioactivity, in comparison with other metallic materials, and have a potential to form CPs on their surface in the living body (Fujibayashi et al., 2003). On the other hand, no CP formation occurs on the surface of Zr due to its high chemical stability (Tsutsumi et al., 2009). This might be an issue to be solved for devices requiring early bone adhesion.
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Ha et al. (2011) investigated the modification of the Zr surface using a micro-arc oxidization (MAO) method (Rocchi et al., 2003). When the MAO treatment for 10 min was performed to Zr in an aqueous solution containing calcium glycerophosphate and magnesium acetate, a porous ZrO2 layer of 5 μm thickness formed on its surface. Subsequently, when the sample was chemically treated with sulfuric acid or sodium hydroxide, some changes at its surface, such as enhancement of oxidation and increase in the amount of hydroxy groups, occurred. When they were soaked in HBSS for 12 d at 37°C, they were completely covered with a thick layer of HCA. On the other hand, Zr treated with only MAO (without the chemical treatments) showed no HCA formation. The biocompatibility of the Zr surface can be improved effectively by combining a MAO method with some chemical treatments.
14.4
Summary
This chapter has briefly reviewed some surface modifications on metallic biomaterials, based on ceramics. Some directions, such as antibacterial hydroxyapatite coatings using thermal spraying and strongly bonded, thin film coatings consisting of a CP phase using rf-magnetron sputtering or MOCVD, have been developed. A bioactive coating using CP glass for harmonizing β-type TNTZ might be one of the most significant items for various biomedical applications. In order to suppress the corrosion of the Mg or its alloys, well-crystallized CP coatings have been prepared by a chemical solution deposition method. Surface modification by alkali-treatment is one of the most useful methods for preparing bioactive surfaces. The “bioactive Ti” has high reliability for long-term use, since the bioactive ceramic phase is integrated with metallic Ti. The bioinert surface of Zr can be converted into a bioactive one by combining a MAO method with chemical treatments. Surface-modification of metallic biomaterials may make new progress for the next generation through the integration with bioresorbable polymers or biological substances such as proteins and/or cells. Although not shown in this chapter, the coatings of polymers on metallic materials are very attractive. Hanawa and his colleagues have reported that segmented polyurethane (SPU) was coated on Ti or TNTZ with a silane coupling agent (mercaptopropyltrimethoxysilane: γ-MPS) (Sakamoto et al., 2007; Hieda et al., 2014). It has been reported that SPU coating could inhibit blood platelet adhesion and biofilm formation (Tanaka et al., 2010) and that immobilization of RGD peptide on SPU could enhance bone formation (Oya et al., 2009). In future, this is expected to develop into metallic biomaterials to which various bio-functional molecules are manipulated.
References Akahori, T., Niinomi, M., Koyanagi, Y., Kasuga, T., Toda, H., Fukui, H., Ogawa, M., 2005. Mechanical properties of biocompatible beta-type titanium alloy coated with calcium phosphate invert glass-ceramic layer. Mater. Trans. 46, 1564–1569.
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Akahori, T., Niinomi, M., Nakai, M., Fukuda, H., Fukui, H., Ogawa, M., 2007. Bioactive ceramic surface modification of β-type Ti-Nb-Ta-Zr system alloy by alkali solution treatment. Mater. Trans. 48, 293–300. Akiyama, T., Miyamoto, H., Yonekura, Y., Tsukamoto, M., Ando, Y., Noda, I., Sonohata, M., Mawatari, M., 2013. Silver oxide-containing hydroxyapatite coating has in vivo antibacterial activity in the rat tibia. J. Orthop. Res. 31, 1195–1200. Allen, G.C., Ciliberto, E., Fragala, I., Spoto, G., 1996. Surface and bulk study of calcium phosphate bioceramics obtained by metal organic chemical vapor deposition. Nucl. Instrum. Methods Phys. Res. B 116, 457–460. Asami, K., Saito, K., Ohtsu, N., Nagata, S., Hanawa, T., 2007. Titanium-implanted CaTiO3 films and their changes in Hanks’ solution. Surf. Interface Anal. 35, 483–488. Berndt, C.C., Haddad, G.N., Farmer, A.J.D., 1990. Thermal spraying for bioceramic applications. Mater. Sci. Forum 14, 161–173. Darr, J.A., Guo, Z.X., Raman, V., Bououdina, M., Rehman, V., 2004. Metal organic chemical vapour deposition (MOCVD) of bone mineral like carbonated hydroxyapatite coatings. Chem. Commun. 62, 455–460. van Dijk, K., Schaeken, H.G., Wolke, J.C.G., Maree, C.H.M., Habraken, F.H.P.M., Verhoeven, J., Jansen, J.A., 1995. Influence of discharge power level on the properties of hydroxyapatite films deposited on Ti-6A1-4V with RF magnetron sputtering. J. Biomed. Mater. Res. 29, 269–276. Ducheyne, P., Raemdonck, W.V., Heughebaert, J.C., Heughebaert, M., 1986. Structural analysis of hydroxyapatite coating on titanium. Biomaterials 7, 97–103. Eto, S., Kawano, S., Someya, S., Miyamoto, H., Sonohata, M., Mawatari, M., 2016. First clinical experience with thermal-sprayed silver oxide-containing hydroxyapatite coating implant. J. Arthroplasty 31, 1498–1503. Fujibayashi, S., Neo, M., Kim, H.M., Kokubo, T., Nakamura, T., 2003. A comparative study between in vivo bone ingrowth and in vitro apatite formation on Na2O-CaO-SiO2 glasses. BiomateriaIs 24, 1349–1356. Ha, J.Y., Tsutsumi, Y., Doi, H., Nomura, N., Kim, K.H., Hanawa, T., 2011. Enhancement of calcium phosphate formation on zirconium by micro-arc oxidation and chemical treatments. Surf. Coat. Technol. 205, 4948–4955. Hashimoto, T., Obata, A., Kasuga, T., 2007. Preparation of silicon-containing apatite coating on titanium. Arch. BioCeram. Res. 7, 133–136 (Proc. Asian BioCeram. 2007). Hench, L.L., 1992. Bioceramics: from concept to clinic. J. Am. Ceram. Soc. 74, 1487–1510. Hieda, J., Niinomi, M., Nakai, M., Cho, K., Mohri, T., Hanawa, T., 2014. Adhesive strength of medical polymer on anodic oxide nanostructures fabricated on biomedical β-type titanium alloy. Mater. Sci. Eng. C 36, 244–252. Hiromoto, S., Yamamoto, A., 2009. High corrosion resistance of magnesium coated with hydroxyapatite directly synthesized in an aqueous solution. Electrochim. Acta 54, 7085–7093. Inagaki, M., Yokogawa, Y., Kameyama, T., 2003. Bond strength and microstructure of radiofrequency thermal plasma sprayed hydroxyapatite/titanium composite coatings. Key Eng. Mater. 240–242, 99–102. Kasuga, T., 2005. Bioactive calcium pyrophosphate glasses and glass-ceramics. Acta Biomater. 1, 55–64. Kasuga, T., Hattori, T., Niinomi, M., 2012. Phosphate glasses and glass-ceramics for biomedical applications. Phosphorous Res. Bull. 26, 008–015. Kato, H., Nakamura, T., Nishiguchi, S., Matsusue, Y., Kobayashi, M., Miyazaki, T., Kim, H.-M., Kokubo, T., 2000. Bonding of alkali- and heat-treated tantalum implant to bone. J. Biomed. Mater. Res. 53, 28–35.
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Kitsugi, T., Yamamuro, T., Nakamura, T., Higashi, S., Kotani, Y., Hyakuna, K., Ito, S., Kokubo, T., Takagi, M., Shibuya, T., 1986. Bone bonding behavior of three kinds of apatite containing glass ceramics. J. Biomed. Mater. Res. 20, 1295–1307. Kokubo, T., Takadama, H., 2006. How useful is SBF in predicting in vivo bone bioactivity? Biomaterials 27, 2907–2915. Kokubo, T., Matsushita, T., Takadama, H., Kizuki, T., 2009. Development of bioactive materials based on surface chemistry. J. Eur. Ceram. Soc. 29, 1267–1274. Kuroda, D., Niinomi, M., Morinaga, M., Kato, Y., Yashiro, T., 1998. Design and mechanical properties of new β type titanium alloys for implant materials. Mater. Sci. Eng. A 243, 244–249. Lacefield, W.R., 1993. Hydroxylapatite coatings (Chapter 12). In: An Introduction to Bioceramics. World Scientific, Singapore, pp. 223–238. Li, T.T., Lee, J.H., Kobayashi, T., Aoki, H., 1996. Hydroxyapatite coating by dipping method, and bone bonding strength. J. Mater. Sci. Mater. Med. 7, 355–357. Li, S.J., Niinomi, M., Akahori, T., Kasuga, T., Yang, R., Hao, Y.L., 2004. Fatigue characteristics of bioactive glass-ceramic-coated Ti–29Nb–13Ta–4.6Zr for biomedical application. Biomaterials 25, 3369–3378. Madi, M., Zakaria, O., Kasugai, S., 2014. Coated vs uncoated implants: bone defect configurations after progressive peri-implantitis in dogs. J. Oral Implantol. 40, 661–669. Madi, M., Zakaria, O., Ichinose, S., Kasugai, S., 2016. Effect of induced periimplantitis on dental implants with and without ultrathin hydroxyapatite coating. Implant. Dent. 25, 39–46. Maeda, H., Tamura, T., Kasuga, T., 2017. Experimental and theoretical investigation of the structural role of titanium oxide in CaO P2O5 TiO2 invert glass. J. Phys. Chem. B 121, 5433–5438. Miyazaki, T., Kim, H.M., Miyaji, F., Kokubo, T., Nakamura, T., 2000. Bioactive tantalum metal prepared by NaOH treatment. J. Biomed. Mater. Res. 50, 35–42. Narushima, T., Ueda, K., Goto, T., Masumoto, H., Katsube, T., Kawamura, H., Ouchi, C., Iguchi, Y., 2005. Preparation of calcium phosphate films by radiofrequency magnetron sputtering. Mater. Trans. 46, 2246–2252. Narushima, T., Ueda, K., Goto, T., Katsube, T., Ouchi, C., Iguchi, Y., 2007. Calcium phosphate films coated on titanium by RF magnetron sputtering for medical applications. Mater. Sci. Forum 539–543, 551–556. Niinomi, M., 2003. Cyto-toxicity and fatigue performance of low rigidity titanium alloy, Ti-29Nb-13Ta-4.6Zr, for biomedical applications and healthcare goods. Biomaterials 24, 2673–2683. Niinomi, M., Hattori, T., Morikawa, K., Kasuga, T., Suzuki, A., Fukui, H., Niwa, S., 2002. Development of low rigidity β-type titanium alloy for biomedical applications. Mater. Trans. 43, 2970–2977. Niinomi, M., Akahori, T., Yamaguchi, T., Kasuga, T., Fukui, H., Suzuki, A., 2003. Ageing characteristics and mechanical properties of Ti-29Nb-13Ta-4.6Zr coated with calcium phosphate glass-ceramic for biomedical applications. J. Jpn. Inst. Metals 67, 604–613. Noda, I., Miyaji, F., Ando, Y., Miyamoto, H., Shimazaki, T., Yonekura, Y., Miyazaki, M., Mawatari, M., Hotokebuchi, T., 2009. Development of novel thermal splayed antibacterial coating and evaluation of release properties of silver ions. J. Biomed. Mater. Res. B Appl. Biomater. 89B, 456–465. Ong, J.L., Lucas, L.C., 1994. Post-deposition heat treatments for ion beam sputter deposited calcium phosphate coatings. Biomaterials 15, 337–341. Oya, K., Tanaka, Y., Saito, H., Kurashima, K., Nogi, K., Tsutsumi, H., Doi, H., Nomura, N., Hanawa, T., 2009. Calcification by MC3T3-E1 cells on RGD peptide immobilized on titanium through electrodeposited PEG. Biomaterials 30, 1281–1286.
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Rocchi, A., Martignori, M., Gottlow, J., 2003. Immediate loading of Bra˚nemark system® TiUniteTM and machined-surface implants in the posterior mandible: a randomized open- ended clinical trial. Clin. lmplant. Dent. Relat. Res. 5, 57–63. Sakamoto, H., Doi, H., Kobayashi, I., Yoneyama, T., Suzuki, S., Hanawa, T., 2007. Structure and strength at the bonding interface of a titanium-segmented polyurethane composite through 3-(trimethoxysilyl) propyl methacrylate for artificial organs. J. Biomed. Mater. Res. A 82, 52–61. Sato, M., Tu, R., Goto, T., Ueda, K., Narushima, T., 2007a. Hydroxyapatite formation on calcium phosphate coated titanium. Mater. Sci. Forum 561–565, 1513–1516. Sato, M., Tu, R., Goto, T., Ueda, K., Narushima, T., 2007b. Hydroxyapatite formation on CaTiO3 film prepared by metal-organic chemical vapor deposition. Mater. Trans. 48, 1505–1510. Tanaka, Y., Matin, K., Gyo, M., Okada, A., Tsutsumi, Y., Doi, H., Nomura, N., Tagami, J., Hanawa, T., 2010. Effects of electrodeposited poly(ethylene glycol) on biofilm adherence to titanbium. J. Biomed. Mater. Res. A 95, 1105–1113. Tomozawa, M., Hiromoto, S., 2011a. Growth mechanism of hydroxyapatite-coatings formed on pure magnesium and corrosion behavior of the coated magnesium. Appl. Surf. Sci. 257, 8253–8257. Tomozawa, M., Hiromoto, S., 2011b. Microstructure of hydroxyapatite- and octacalcium phosphate-coatings formed on magnesium by a hydrothermal treatment at various pH values. Acta Mater. 59, 355–363. Tsui, Y.C., Doyle, C., Clyne, T.W., 1998. Plasma sprayed hydroxyapatite coatings on titanium substrates, part 1: mechanical properties and residual stress levels. Biomaterials 19, 2015–2029. Tsutsumi, Y., Nishimura, D., Doi, H., Nomura, N., Hanawa, T., 2009. Difference in surface reactions between titanium and zirconium in Hanks’ solution to elucidate mechanism of calcium phosphate formation on titanium using XPS and cathodic polarization. Mater. Sci. Eng. C 29, 1702–1708. Tsutsumi, H., Niinomi, M., Nakai, M., Gozawa, T., Akahori, T., Saito, K., Tu, R., Goto, T., 2010. Fabrication of hydroxyapatite film on Ti-29Nb-13Ta-4.6Zr using a MOCVD technique. Mater. Trans. 51, 2277–2283. Ueda, K., Narushima, T., Goto, T., Katsube, T., Nakagawa, H., Kawamura, H., Katsube, T., 2007. Evaluation of calcium phosphate coating films on titanium fabricated using RF magnetron sputtering. Mater. Trans. 48, 307–312. Wolke, J.G.C., JPCM, v.d.W., Schaeken, H.G., Jansen, J.A., 2003. In vivo dissolution behavior of various RF magnetron-sputtered Ca-P coatings on roughened titanium implants. Biomaterials 24, 2623–2629. Yoshinari, M., Hayakawa, T., Wolke, J.G.C., Nemoto, K., Jansen, J.A., 1997. Influence of rapid heating with infrared radiation on RF magnetron-sputtered calcium phosphate coatings. J. Biomed. Mater. Res. 37, 60–67.
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Toshihiro Kasuga Division of Advanced Ceramics, Nagoya Institute of Technology, Nagoya, Japan
14.1
Introduction
The high-performance biomaterials for replacing damaged hard tissues, such as artificial joints, bone fillers, dental implants, and so on, are in high demand among the aged people. Metallic biomaterials with biological safety, such as stainless steel, cobalt (Co)-based alloys, titanium (Ti) and its alloys, are reasonable for replacing hard tissues at load-bearing sections. Ti and its alloys are especially widely used as orthopedic and dental implants because of their excellent biocompatibility. Much attention has been paid to β-type Ti alloys because they have high mechanical strength with low modulus of elasticity. One of the attractive metals is β-type Ti alloys, which contain elements sufficiently tolerated by living tissues, such as niobium (Nb), tantalum (Ta), and zirconium (Zr) (Niinomi, 2003; Kuroda et al., 1998). A Ti-29Nb-13Ta-4.6Zr (denoted by TNTZ) alloy, which has been suggested as one of the best materials designed, shows a high tensile strength of 1 GPa and low Young’s modulus of 60–80 GPa. Although Ti and its alloys do not make a chemical bond with living bone, they can adhere to it with almost no fibrous tissues. Calcium phosphate ceramics such as hydroxyapatite (Ca10(PO4)6(OH)2: HA) and β-tricalcium phosphate (β-Ca3(PO4)2: β-TCP) exhibit bioactivity, which is the ability to make a bond chemically with natural bone tissues (Hench, 1992). Some glasses and glass-ceramics are also well known as bioactive materials. Some portions in artificial substitutes for hard tissue may require bioactivity. Bioactive surface modification is, in general, applied to Ti and its alloys to further improve their bio-functionality, including osteoconductivity. The bioactive ceramic coatings prepared using various processes have been investigated so far and some materials are clinically used as biomaterials for dental and surgical implants. In this chapter, some bio-functional coatings on metals are briefly reviewed.
14.2
Calcium phosphate ceramic coatings
Implants coated with calcium phosphates such as HA and β-TCP have been widely used in orthopedics and dentistry. The cementless fixation technique combines the mechanical strength, ductility, and ease of fabrication with bioactivity. Since much earlier, calcium phosphate coatings have been applied by a variety of methods, such Metals for Biomedical Devices. https://doi.org/10.1016/B978-0-08-102666-3.00014-6 © 2019 Elsevier Ltd. All rights reserved.
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as dip-coating (Li et al., 1996), electrophoretic deposition (EPD) (Ducheyne et al., 1986), plasma spraying (Lacefield, 1993), ion-beam sputtering (Ong and Lucas, 1994), flame spraying (Berndt et al., 1990), wet-process coating, and so on. In this section, some recent possible coatings are briefly discussed.
14.2.1 Thermal sprayed coatings Thermal spraying employs a plasma/ionized gas or fuel/oxygen (O) gas to melt the ceramic particles and subsequently to carry them onto the surface of substrates. Two important properties of thermal sprayed coatings are: (i) bonding strength between the calcium phosphate layer and the metallic substrate, and (ii) dissolution behavior in body fluid. Residual stress arises at the interface between the metallic substrate and the coating layer due to the large difference in their thermal expansions. Amorphous calcium phosphate tends to dissolve rapidly in body fluid: coatings with low crystallinity rapidly become weak and may induce inflammatory responses. A relatively high bonding strength on Ti or its alloys was reported by Tsui et al. (1998); the values were in the range of 20–40 MPa after optimization of coating properties. The crystallinity of HA in their coating layers was relatively high (83%–94%). Their high crystallinity was very important for the long-term benefit of the coated implants. A radio-frequency thermal plasma spraying (rf-TPS) method was developed to obtain strong bonding between the HA coating and Ti substrate, employing an HA/ Ti composite coating (Inagaki et al., 2003). The HA/Ti composite coatings were prepared by controlling the starting composition using two microfeeders, which fed the HA and Ti powders at an accurate rate. The starting powders were fed so as to form the coated layer with a compositional gradient from a Ti-rich phase at the bottom to an HA-rich at the top. The composition of plasma gas played an important role in improving the bonding strength of the coatings. Rf-TPS with nitrogen (N)-containing plasma gas induced the formation of the coatings with excellent bonding to the Ti substrate (the tensile bonding strength of 65 MPa). The microstructure of the interface between deposited Ti particles in the coatings was significantly influenced by plasma gas composition during rf-TPS. By rf-TPS with N- or O-containing plasma gas, Ti nitride or Ti oxide was formed in the deposited Ti particles, respectively. The bonding of the composite coatings on the Ti substrate was suggested to be influenced by microstructure at the interface between Ti particles. Surgical Site Infection (SSI) is one of the complications in artificial joint replacement surgery, and its treatment is often not easy. About 50% of the infecting organisms are Methichillin-resistant Staphylococcus aureus (MRSA) and methicillinresistant Staphylococcus epidermidis (MRSE). The SSI associated with the surgery, which is called “implant related infection” or “peri-prosthetic joint infection”, is known to be linked to the formation of biofilm and it has a serious influence on the outbreak and severity of infectious disease. Antibacterial drugs don’t work well on bacteria in biofilms, and the treatment of SSI is difficult due to an increase in drug-resistant bacteria such as MRSA and MRSE. A silver (Ag)-containing calcium phosphate (CP) coating has been prepared using frame spraying to reduce the incidence of SSI in total hip arthroplasty (THA) (Noda et al., 2009). A mixture of 97% HA and 3% Ag2O powders (in wt%) was
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flame-sprayed on metallic Ti using an acetylene torch to be coated with a compositionally homogeneous amorphous calcium phosphate layer of 40 μm thickness (denoted by Ag-CP coating, hereafter). As shown in Fig. 14.1A and B, a typical surface structure of thermal spraying was obtained. Energy dispersive x-ray spectrometry (EDX) showed that the content of Ag2O in the sprayed layer was reduced to 0.3%. However, its calcium (Ca)/phosphate (P) ratio was almost unchanged from the nominal composition. The amount of Ag+ ions dissolved from the Ag-CP coating in fetal bovine serum (FBS) was 78 ppb after 1 h of soaking and 260–330 ppb after 24–48 h, respectively; Ag+ ion dissolution at an initial stage would be important for preventing SSI (Noda et al., 2009).
Fig. 14.1 Scanning electron microscopic (SEM) images of (A) surface and (B) cross-section of Ag-CP coating. No silver-related particles were observed. Reprinted from Noda, I., Miyaji, F., Ando, Y., Miyamoto, H., Shimazaki, T., Yonekura, Y., Miyazaki, M., Mawatari, M., Hotokebuchi, T., 2009. Development of novel thermal splayed antibacterial coating and evaluation of release properties of silver ions. J. Biomed. Mater. Res. B Appl. Biomater. 89B, 456–465. Copyright 2008, with permission from Wiley.
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After phosphate buffer solution (PBS) containing 102 CFU of MRSA was injected in a hole drilled into the tibial medullary cavity of 10-week-old Sprague-Dawley rats, Ag-CP coating and conventional CP flame-sprayed coating (denoted by CP coating, hereafter) samples were implanted and closed. After 24, 48, and 72 h, the tibia including the implant was dissected out and then crushed and ultrasonically irrigated with PBS. The dilution of this PBS was incubated on agar plates at 37°C and the number of inoculated viable cells was calculated (a dilution plate method). It was confirmed that the number of bacteria in the case of Ag-CP coating was significantly smaller (P < .02) than that in the case of CP coating (Akiyama et al., 2013). The silver amounts in the serum of artificial joint replacement surgery using Ag-CP coating showed a maximum value at day 14 postoperatively (average value 1.4 ng/mL), and then they gradually decreased. During the experimental period, no troubles caused by argyria were observed; high safety was indicated (Eto et al., 2016). Osteoconductivity of the Ag-CP and CP coatings was evaluated in tibia models using rats and rabbits. Newly formed bone was observed on the surface layer of both the samples, and there was no significant difference between them. Both of them were concluded to have excellent osteoconductivity. The Ag-CP coating is emphasized to be a great coating technology for the formation of functional surface with both osteoconductivity and antibacterial activity. The coating has obtained the clearance of Japanese Ministry of Health, Labor and Welfare in 2016 for femoral stem and acetabular cup of THA.
14.2.2 Thin film coatings Physical vapor deposition (PVD) is one of the most promising methods to obtain thin, uniform, and dense CP coatings on metallic substrates. Radiofrequency (rf ) magnetron sputtering has been used widely for coatings of thin films with excellent bonding to the substrates and has been applied to the coatings of CP films on commercially pure Ti (cpTi) (Yoshinari et al., 1997; Wolke et al., 2003) and Ti alloys (van Dijk et al., 1995), and the method has a great advantage in that the coatings can be achieved at relatively low temperatures. That is, no degradation in the mechanical properties of the metallic substrates occurs. Narushima et al. (2005) prepared CP films on cpTi substrates by rf magnetron sputtering using a hot-pressed, dense β-TCP target. The resulting films were dense and uniform with a smooth surface, consisting of amorphous CP and/or oxyapatite phases (Fig. 14.2; Ueda et al., 2007). The bonding strengths between the coating and the cpTi substrate were estimated to be >60 MPa. When the film was soaked in Hank’s balanced salt solution (HBSS) or PBS, the dissolved amount of Ca2+ ions from amorphous coatings was significantly larger than that from oxyapatite coatings. Animal test using beagle dogs showed that cpTi coated with the CP film was fixed rapidly and strongly with bone (Narushima et al., 2007; Ueda et al., 2007). CpTi with a sputtered HA thin layer was developed as dental implants in Japan. Animal tests using canine mandibles for 28 weeks showed excellent bone-to-implant contact (BIC) by scanning electron microscopic analysis: the values of HA-sputtered implants showed larger (98.1%), in comparison with those of machined implants
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Fig. 14.2 Cross-sectional SEM image of calcium phosphate film on a cpTi plate prepared by rf magnetron sputtering. Reprinted from Ueda, K., Narushima, T., Goto, T., Katsube, T., Nakagawa, H., Kawamura, H., Katsube, T., 2007. Evaluation of calcium phosphate coating films on titanium fabricated using RF magnetron sputtering. Mater. Trans. 48, 307–312. Copyright 2007, with permission from The Japan Institute of Metals.
(without HA-coatings) (70.4%) (Madi et al., 2016). The HA coating layer was almost completely dissolved. By evaluating the type and dimensions of bone defects after progressive peri-implantitis in dogs, peri-implant defects around the HA-sputtered implants were smaller than those around plasma-sprayed HA-coated implants (Madi et al., 2014). The dental implant system using HA-sputtered cpTi has been commercialized in Japan since 2010. Chemical vapor deposition (CVD) also has great advantages in the preparation of CP thin films coated strongly onto metallic substrates, because of high deposition rates with excellent microstructure controllability. Some reports have been published on the preparation of β-Ca2P2O7 film (Allen et al., 1996) and its conversion to β-TCP phase by postheat treatment. A few works on the direct preparation of bioactive CP films such as HA or β-TCP have been reported so far (Darr et al., 2004; Sato et al., 2007a). Bioactive CP films have been prepared on cpTi substrates by metal-organic chemical vapor deposition (MOCVD) using bis-dipivaloylmetanato-calcium (Ca(dpm)2) and tryphenyl phosphate ((C6H5O)3PO) precursors (Sato et al., 2007a). A dense HA film was prepared by controlling Ca/P atomic ratio and substrate temperature at 700–800°C. The resulting HA film showed (002) preferred orientation parallel to the substrate surface. The film can be deposited at high rates of 4–6 nm/s. HA formed in HBSS at 37°C within 6 h on the HA film. The film on cpTi was expected to show excellent bioactivity. Tsutsumi et al. (2010) prepared a polycrystalline HA film on a TNTZ surface using MOCVD by controlling the heating temperature, as shown in Fig. 14.3. The thickness of the HA film on TNTZ could be varied by 1–7 μm with the deposition time. An α-phase precipitated in the TNTZ matrix after heating the substrate during the deposition process, and the mechanical properties of HA-coated TNTZ were improved by maintaining its low Young’s modulus. MOCVD would be one of the most effective methods for HA coating on TNTZ. Calcium titanate (CaTiO3) film has also been investigated as one of the promising bioactive coatings, using a sputtering (Asami et al., 2007) or CVD (Sato et al., 2007b)
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Fig. 14.3 SEM images of surfaces of (A) TNTZ, (B) surface, and (C) cross-section of HA coated on TNTZ using MOCVD. HA deposited on TNTZ had a dense granular microstructure and its grain diameter was 2 μm. Reprinted from Tsutsumi, H., Niinomi, M., Nakai, M., Gozawa, T., Akahori, T., Saito, K., Tu, R., Goto, T., 2010, Fabrication of hydroxyapatite film on Ti-29Nb-13Ta-4.6Zr using a MOCVD technique, Mater. Trans. 51, 2277–2283. Copyright 2010, with permission from The Japan Institute of Metals.
method. HA formed on the film after soaking in HBSS for 3–7 days. It was proposed that the CaTiO3 film would firstly dissolve into the solution to locally increase Ca2+ ion concentration, where PO34 ions could be attracted and reacted with Ca2+ ions, resulting in the formation of HA.
14.2.3 Apatite coatings using wet-processing At the interface between bioactive ceramics and bone in living body, an apatite layer can be often seen: the ceramics bond to living bone through the apatite layer (Kitsugi et al., 1986), which consists of a hydroxycarbonate apatite (HCA) with small crystallites and defective structure, precipitated by incorporating carbonate anions from body fluid within amorphous CP phase. HCA has a similar composition and structure to bone. The apatite layer can form on the ceramics, metals, and polymers when they are soaked in simulated body fluid (SBF) with ion concentrations nearly equal to those of human blood plasma (Kokubo and Takadama, 2006). Two indispensable conditions needed for the HCA formation on materials using SBF are: (i) the existence of surface functional groups that induce the nucleation of apatite (e.g., Si-OH, Ti-OH, Nb-OH, Ta-OH, COOH, and PO4H2 groups), and (ii) the increase in the supersaturation concerning apatite around the surface of the materials. Metallic Ti and its alloys are generally covered with a passive layer of Ti oxide. Since the surface of the layer includes the Ti-OH group, there exists the possibility of HCA formation. However, relatively high supersaturation concerning HCA is usually needed around the surface. To increase the supersaturation in SBF, a large amount of Ca2+ ions should be supplied. For example, soluble Ca carbonate (vaterite) was expected to supply both Ca2+ and carbonate ions when soaked in SBF: the supersaturation concerning HCA increased in SBF and rapid HCA-formation occurred (Hashimoto et al., 2007). However, the mechanically adhesive strength of the HCA layer on metallic substrates might be weak.
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Fig. 14.4 Cross-sectional SEM and element distribution images of HA-coating on Mg prepared by a chemical solution deposition method at pH 8.9. The sample was scratched off from the substrate. A thin Mg(OH)2 layer formed at the bottom of the coating. The intensity of Ca and P distribution was almost constant in the outer HA-layer packed densely. Reprinted from Tomozawa, M., Hiromoto, S., 2011b. Microstructure of hydroxyapatite- and octacalcium phosphate-coatings formed on magnesium by a hydrothermal treatment at various pH values. Acta Mater. 59, 355–363. Copyright 2010, with permission from Elsevier.
Metallic Mg or its alloys have been investigated as a biodegradable material for applications in medical implants such as vascular stents. In application to orthopedic devices such as fracture fixation materials and artificial bones, there exists a serious issue with Mg corrosion, which proceeds faster than healing of the diseased area. Therefore, in order to suppress the corrosion of the Mg or its alloys, a chemical solution deposition method for developing well-crystallized CP coatings on them, has been reported (Hiromoto and Yamamoto, 2009). An HA coating was prepared by soaking in C10H12CaN2-Na2O8 (Ca-EDTA) KH2PO4 (their concentration: 0.25 mol L 1) adjusted to pH 9–12 at 90°C for 2 h. The coating showed a double-layer structure, consisting of a continuous inner layer of 1 μm thickness and an outer layer of rod-shaped HA crystals grown on the inner layer to a length of 4 μm, as shown in Fig. 14.4. Almost no nanopores were observed in the inner layer, and the coating was dense at its micrometer level (Tomozawa and Hiromoto, 2011a). From the anodic polarization curve of HA-coated or noncoated pure Mg in 3.5 mass% NaCl aq., it was shown that the cathode current density of the coated sample was much smaller than that of the noncoated one (Tomozawa and Hiromoto, 2011b). It has been shown that the corrosion of Mg or its alloys could be suppressed drastically by the dense HA coating.
14.2.4 Calcium phosphate glass-ceramic coatings Since glass-based materials can be allowed some latitude in choice of composition, they have a significant advantage in that they are controllable to improve their physico-chemical properties and microstructures after their crystallization to attain sufficient performance as biomaterials. Silicate-based glass and glass-ceramics such as Bioglass and Cerabone A-W are well-known as bioactive materials (Hench, 1992).
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In some types of glass, anionic groups can be connected through a cation, which should usually be a network modifier, to form the glassy state; such glass types are called “invert glasses”. Kasuga (2005) reported silicate-free invert glass with compositions of high CaO and low P2O5 contents; the glass consisted of PO34 (orthophosphate) and/or P2O47 (pyrophosphate) groups without PO3 (metaphosphate) group, and the phosphate groups were connected through Ca2+ ions. The proposed mother composition was 60CaO-30P2O5-3TiO2-7Na2O (in mol%, denoted by PIG hereafter). It is usually difficult to obtain glass with Ca/P > 0.75 in the binary system, but they can be prepared with small amounts of TiO2 and Na2O. TiO2 addition into phosphate glass improved their glass-forming ability and chemical durability. The structural role of TiO2 in the CP invert glass was clarified by spectroscopic analysis and molecular dynamics simulation (Maeda et al., 2017). PIG was pulverized to <10 μm in diameter (the average size; 1.0–1.5 μm) The glass powder slurry was dip-coated onto a sandblasted TNTZ and then dried. Then the substrate with the glass powder layer was heated in air at 800°C for 1 h and cooled to room temperature. Due to sintering of the glass powder, a layer with 10 μm thickness was formed on TNTZ. Many pores of several micrometers in size were seen. There existed a reaction zone of 2–4 μm in thickness between the glass-ceramic layer and the substrate. The reaction layer was spontaneously developed on TNTZ by heating in air to relax the thermal stress generated in the coating layer and the substrate, resulting in the formation of the strong joint. The coating layer in the sample consisted predominantly of β-Ca3(PO4)2 (β-TCP) phase with trace amounts of β-Ca2P2O7 (β-CPP) and TiO2 (rutile phase) crystalline phases. The tensile bonding strength of the glass-ceramic layer on TNTZ was estimated to be 26 MPa. The fracture occurs in the glass-ceramic layer with porosity; the bonding strength at the interface layer between TNTZ and the glass-ceramic was higher than the tensile strength of the glass-ceramic (Kasuga et al., 2012). The oxidized layer at the TNTZ surface was estimated to be <0.5 μm in depth. This thin oxidized layer worked to form the strong bond to PIG glass-ceramic. Aging treatment of TNTZ at 400°C is effective for increasing its hardness, modulus of elasticity, and strength (Niinomi et al., 2002). This improvement is due to the precipitation of ω-phases in the matrix β-phase during the aging. After the glassceramic-coated TNTZ was heat-treated for aging, tensile bonding strength of the sample with the coating layer of 5 μm thickness was not degraded, while that of 20 μm thickness was degraded due to the generation of microcracks, which originated from the formation of ω-phase, at the interface between the layer and the substrate (Niinomi et al., 2003; Akahori et al., 2005). The coated samples were machined for fatigue tests according to ISO1099-75 specifications (Li et al., 2004). No delaminations of the coating were observed during the fatigue test. The fatigue resistance of the samples was improved due to an increase in resistance to fatigue crack initiation caused by the formation of the α-phase during heating for the coating and the precipitation of the ω-phase from the substrate during the slow cooling. In addition, aging at 400°C for 72 h greatly increased fatigue limit of the coated TNTZ and effectively relieved the detrimental effect of sandblasting on fatigue properties.
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Fig. 14.5 Contact microradiographs around the glass-ceramic-coated TNTZ rods (5 mmϕ) implanted into femurs of Japanese rabbits after 1 month, 1 year, and 5 years. The scale bar indicates 50 μm. Reprinted from Kasuga, T., Hattori, T., Niinomi, M., 2012. Phosphate glasses and glassceramics for biomedical applications. Phosphorous Res. Bull. 26, 008–015. Copyright 2012, with permission from Japanese Association of Inorganic Phosphorus Chemistry.
A rod-shaped sample of 5 mm diameter 10 mm length was implanted into the femur of a Japanese rabbit. The sample was autoclaved in water at 121°C for 1 h. Fig. 14.5 shows contact microradiographs (CMR) after 1 month, 1 year, and 5 years of implantation. No significant change in the thickness of the glass-ceramic coating layer after the implantation could be seen. The observation showed new bone formation around the autoclaved glass-ceramic-coated TNTZ after 1 month of implantation and the bone tissue showed direct contact with implants. After 1 year and 5 years of implantation, the glass-ceramic-coating was surrounded by maturated bone tissue and had bonded directly with them. When the samples were sliced and polished for the observation after 1 year of implantation, no crack occurred between the coating and bone and also between the coating and TNTZ. This result shows excellent bioactivity of the PIG glass-ceramic coating on TNTZ (Kasuga et al., 2012).
14.3
Bioactive surface prepared by chemical treatments
Bioactive surface modifications described in this section might not be included in the category of coatings on metallic substrates. However, they are briefly reviewed, since they are very important, practical techniques for preparing bioactive surfaces on metallic substrates.
14.3.1 Chemical treatment with a concentrated alkaline solution Ti and its alloys react with alkaline aqueous solutions. The reacted surface consists of negatively charged HTiO3 nH2O, which incorporates alkaline ions. Kokubo et al. prepared amorphous sodium titanate on cpTi by soaking the metal in an aqueous NaOH l
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solution of 5 M at 60°C for 24 h and subsequent heating at 600°C for 1 h. A surface consisting of a porous layer was obtained. The sodium titanate phase was integrated with the metal substrate through a graded structure within a depth of 1.5 μm from the surface. When the treated sample was soaked in SBF, HCA formed on the porous sodium titanate layer. Since the HCA showed a graded structure, it achieved a tight adhesion to the metal substrate (tensile bonding strength; >30 MPa). When Ti alloys, such as Ti-6Al-4V, Ti-6Al-2Nb-Ta and Ti-15Mo-5Zr-3Al, were treated with NaOH and subsequent heating, they also formed sodium titanate phase, which was free of the alloying components. These treatments had no negative impact on the strength under tensile loading nor fatigue strength under cyclic loading of the metals in saline solution (Kokubo et al., 2009). The treated Ti metal and its alloys have been investigated in some animal models. When a cylindrical rod of the treated cpTi was implanted into the intramedullar canal of a rabbit femur, apatite formed on its surface within 3 weeks and it was completely covered with bone within 12 weeks. The rectangular sample implanted into a rabbit tibia formed an HCA layer on its surface and bonded to bone through the layer after 8 weeks. Surface modified Ti and its alloys started to be clinically used as artificial hip joints since 2007. Modification of a TNTZ surface using this type of treatment was also examined; HCA-forming ability was lower on the surface of alkali-treated TNTZ than that on the alkali-treated cp-Ti (Akahori et al., 2007). The HCA-forming ability might be affected due to the existence of the sodium niobate film, which was considered to delay the exchange between the Na+ and H3O+ ions. The surface modification technique to induce bioactivity can also be applied to tantalum (Ta) metal with excellent fracture toughness and malleability (Miyazaki et al., 2000). Typical treatment conditions were subjected to NaOH aq of 0.5 M at 60°C for 24 h and subsequently heated at 300°C. After the treatment, amorphous sodium tantalate phase formed on its surface. Since Na+ ions in the phase on Ta were exchanged with the H3O+ ions in SBF to form Ta-OH groups for inducing apatite nucleation, HCA formed on the treated Ta. This metal also formed apatite in the living body, and bonded tightly to bone through the apatite layer (Kato et al., 2000).
14.3.2 Surface modification by micro-arc oxidation and chemical treatment Since metallic Zr and its alloys show excellent physico-chemical properties similar to Ti and small artifact effect in MRI, they are expected to perform well as new biomaterials. It is known that Ti and its alloys show high biocompatibility close to bioactivity, in comparison with other metallic materials, and have a potential to form CPs on their surface in the living body (Fujibayashi et al., 2003). On the other hand, no CP formation occurs on the surface of Zr due to its high chemical stability (Tsutsumi et al., 2009). This might be an issue to be solved for devices requiring early bone adhesion.
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Ha et al. (2011) investigated the modification of the Zr surface using a micro-arc oxidization (MAO) method (Rocchi et al., 2003). When the MAO treatment for 10 min was performed to Zr in an aqueous solution containing calcium glycerophosphate and magnesium acetate, a porous ZrO2 layer of 5 μm thickness formed on its surface. Subsequently, when the sample was chemically treated with sulfuric acid or sodium hydroxide, some changes at its surface, such as enhancement of oxidation and increase in the amount of hydroxy groups, occurred. When they were soaked in HBSS for 12 d at 37°C, they were completely covered with a thick layer of HCA. On the other hand, Zr treated with only MAO (without the chemical treatments) showed no HCA formation. The biocompatibility of the Zr surface can be improved effectively by combining a MAO method with some chemical treatments.
14.4
Summary
This chapter has briefly reviewed some surface modifications on metallic biomaterials, based on ceramics. Some directions, such as antibacterial hydroxyapatite coatings using thermal spraying and strongly bonded, thin film coatings consisting of a CP phase using rf-magnetron sputtering or MOCVD, have been developed. A bioactive coating using CP glass for harmonizing β-type TNTZ might be one of the most significant items for various biomedical applications. In order to suppress the corrosion of the Mg or its alloys, well-crystallized CP coatings have been prepared by a chemical solution deposition method. Surface modification by alkali-treatment is one of the most useful methods for preparing bioactive surfaces. The “bioactive Ti” has high reliability for long-term use, since the bioactive ceramic phase is integrated with metallic Ti. The bioinert surface of Zr can be converted into a bioactive one by combining a MAO method with chemical treatments. Surface-modification of metallic biomaterials may make new progress for the next generation through the integration with bioresorbable polymers or biological substances such as proteins and/or cells. Although not shown in this chapter, the coatings of polymers on metallic materials are very attractive. Hanawa and his colleagues have reported that segmented polyurethane (SPU) was coated on Ti or TNTZ with a silane coupling agent (mercaptopropyltrimethoxysilane: γ-MPS) (Sakamoto et al., 2007; Hieda et al., 2014). It has been reported that SPU coating could inhibit blood platelet adhesion and biofilm formation (Tanaka et al., 2010) and that immobilization of RGD peptide on SPU could enhance bone formation (Oya et al., 2009). In future, this is expected to develop into metallic biomaterials to which various bio-functional molecules are manipulated.
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