Fabrication of a cubic zirconia nanocoating on a titanium dental implant with excellent adhesion, hardness and biocompatibility

Abstract

A crystalline cubic zirconia (ZrO₂) nanocoating was fabricated in situ on commercially pure titanium metal (cpTi) as a superior dental implant with enhanced biocompatibility. The crystallinity of the nanocoating was achieved at a moderately lower annealing temperature (350 °C) in air using a simple sol–gel method applying a successive layer-by-layer dip-coating technique. Such a procedure facilitates the growth of cubic phase zirconia without noticeable deterioration of the metallic Ti. The bioactivity and biocompatibility of this ZrO₂ coated cpTi (Z-cpTi) were assessed in terms of apatite precipitation through immersion in simulated body fluid, and cell proliferation, respectively for several periods of time. The newly designed Z-cpTi showed unique apatite forming ability, better biocompatibility and tissue attachment properties, and is expected to reduce inflammatory response compared to the bare Ti implants. Moreover, the chemical inertness, corrosion and wear resistant properties, high strength, and appearance of this ZrO₂ nanocoating suggest the probable expediency of Z-cpTi as an advanced oral implant for long-standing performance.

---

Introduction

Metallic implants are scientifically accepted and widely used as scaffolds for load-bearing or bone-contacting applications such as joint and tooth replacement, fracture healing, and renovation of congenital skeletal abnormalities.^1–3 Among them, commercially pure titanium (cpTi) and its alloys are well established and have become the gold standard as oral implants for tooth substitutes due to their potential immunologic and promising aesthetic compromises.^2,4–6 Though Ti based implants demonstrate standard biocompatibility and mechanical properties, they are not enough to construct the implant worthy and durable for longer period of time in corrosive body environment like body fluid.¹ Furthermore, Ti has certain disadvantages like inflammatory response, allergic reactions and galvanic side effects after coming contact with saliva and fluoride.^2,4 Therefore, surface modification by generating a protective as well as bioactive coating is thought to be essential for the improvement of sustainability and suitability of those Ti implants in physiological environment.^1,2 Applying bioactive hydroxyapatite (HA) coating on the Ti surface, the bone bonding ability of the implant could be activated. However, there are a number of disadvantages of a normal HA coating on Ti implant especially in case of dental/oral application. Typically cracks, poor densification and lack of adhesion can be observed in such HA coating on metallic implant due to the thermal stress or irregular interparticle spacing problems.⁷ Furthermore, the HA coating is not chemically inert to sustain in saliva and fluoride containing corrosive oral environment. Therefore, poor mechanical property and lack of chemical inertness of this HA coating can create serious problems after implantation such as risk of delamination, fast dissolution, corrosion, deficiency in osteoconductivity and implant-associated infections due to bacterial colonization.^8,9

It is noteworthy here that zirconia (ZrO₂) is one of the productive ceramic materials for biomedical applications because of its appearance, fracture toughness, chemical inertness, thermal stability and bone-bonding capability.^{2–4,10–17} Pure ZrO₂ is allotropic and polymorphic at ambient temperature and pressure exhibiting three different crystallographic shapes cubic, tetragonal and monoclinic.^12,13,18,19 Out of these, cubic ZrO₂ is the most stable form and recognized for its constructional, and functional properties with elevated mechanical strength, wear resistance, prevention of crack propagation ability, thermal insulation, erosion and alkali resistance.¹⁸ In addition, it demonstrates excellent biocompatibility,²⁰ non-cytotoxicity,^6,19 and moderate bioactivity to promote bonding between the implant and surrounding living tissues.^{10,11,21–23} ZrO₂ has been also shown to enhance osseoconductivity of the implant and reduce bacterial colonization, thus becoming very interesting materials in prosthetic dentistry.^4,17,24,25 Therefore, depositing cubic ZrO₂ as nanocoating on the surface of Ti could improve the anti-corrosion, anti-wear activities, biocompatibility and osseointegrativity of the dental implants as well as minimize inflammatory response of the oral replacements.^2,26 Few groups have developed pure ZrO₂ or ZrO₂ based composite coatings on Ti/Ti-alloy implants by means of magnetron sputtering and plasma spray deposition.^8,16,24,27 However, most of the cases those coatings are thick, inhomogeneous and suffer from cracks and low adhesion.^16,27 Further, necessary biological characterizations of those coatings were also not reported. Nevertheless, there are a number of complications in developing such crystalline uniform coating on Ti metal surface with proper hardness and adhesion. To achieve the enhanced mechanical property, desired phase and crystallinity, the ZrO₂ nanocrystals should be grown in situ on the metal surface. For this purpose, the coated samples must be heat-treated at a relatively high temperature,^12,13 and during such annealing process Ti can be oxidised, contaminated or thermally deteriorated which may depreciate its original chemical properties, biological responses and mechanical strength.

Keeping in view all the issues and probable necessities of chemically inert crystalline ZrO₂ to design an efficient advanced oral implant, we fabricated a thin ZrO₂ coating on cpTi plates by simple sol–gel process applying layer-by-layer dip-coating technique with systematic thermal treatment. The fine crystalline cubic ZrO₂ nanocoating on cpTi (Z-cpTi) with excellent adhesion hardness and homogeneity has been achieved in situ at reasonably low annealing temperature of 350 °C without sacrificing the intrinsic properties of metallic Ti. This type of crystalline cubic ZrO₂ nanocoating on Ti metal with better hardness, appreciable biocompatibility and bioactivity developed by such attractive and simple synthetic route, has not been reported before as the probable candidate for superior dental implant. The fabricated nanocoating was characterized by XRD, FTIR, FESEM and TEM. The primary cellular response and apatite forming ability of Z-cpTi were investigated in details.

Experimental

Materials and methods

All chemicals were used as received. Zirconium(IV) n-propoxide (ZP) in n-propanol (70%) was purchased from Alfa-aesar. 1-Propanol and nitric acid were purchased from RANKEM. 1-Butanol and acetylacetone (acac) were bought from Merck. Water used for hydrolysis was obtained from Milli-Q System (Millipore, Bangalore, India). Commercially pure titanium (grade 4, 99% pure) as metallic dental implants was used for coating purpose.

Preparation of 5 wt% ZrO₂ sol for coating purpose

At first 9.4 g ZP was dissolved in 9.4 g 1-butanol in a beaker. Simultaneously, in another 100 mL beaker 1.4 g acetylacetone and 9 g 1-butanol were taken. The ZP solution was then added into the acetylacetone solution with stirring, and the stirring was continued for 1 h for complete chelation to stabilize the ZP.^28,29 The molar ratio of ZP : acac was 2 : 1. After that 1 g distilled water, 5 g 1-propanol and 0.02 g 1(N) HNO₃ were added and mixed well with the stabilized ZP solution by constant stirring. At last 5 g 1-propanol and 10 g 1-butanol mixture was added to the solution, and stirred at room temperature for 1 h to prepare 5 equivalent wt% ZrO₂ sol. The whole solution was kept overnight in the refrigerator prior to deposition of the coatings.

The final ZrO₂ sol was used to fabricate nanocoating on cpTi plates by applying dip-coating technique (Dip-master 200, Chemat Corporation), adjusting withdrawal velocity to 13 cm min⁻¹. Initially, the as prepared coatings were dried at 70 °C for 1 h and then the single layer coated plates were again dip-coated two times by following the similar procedure to get final (three layer) ZrO₂ coating on cpTi metal plates (designated as “Z-cpTi”). Finally, the Z-cpTi samples were heat-treated at 350 °C in air with a ramp of 2 °C min⁻¹ for 10 h in a SiO₂ glass tube furnace. It should be noted here that consecutive deposition of three layers along with systematic thermal treatment was necessary for the growth of sufficient crystallinity in cubic phase of ZrO₂ at such low annealing temperature. Concurrently, this three layer ZrO₂ coating was also deposited on glass and intrinsic Si wafer under identical condition as representative samples for some characterization purposes.

Characterizations

Grazing incident X-ray diffraction (GIXRD) pattern of the fabricated Z-cpTi was recorded using a Rigaku SmartLab X-ray diffractometer operating at 9 kW (200 mA; 45 kV) using Cu Kα (λ = 1.54059 Å) radiation in thin film mode at a scan rate 2° 2θ s⁻¹. A grazing incidence angle of 0.3° was maintained to record the XRD pattern of the coatings. Materials Data, Inc. (MDI) Jade 7 program and the references of international centre for diffraction data (ICDD) were used to evaluate the XRD data of the samples. Fourier transformed infrared (FTIR) absorption spectra of the dried and final heat-treated coatings (deposited on intrinsic Si wafers) were recorded using Nicolet 380 FTIR spectrophotometer with 512 scans. Transmission electron microscopic (TEM) measurements were carried out with Tecnai G2-30ST (FEI) operating at 300 kV attached with EDX facility. A small amount of scratched off film from coated glass was placed on a carbon coated Cu-grid. The thickness and refractive index (n) of the ZrO₂ films (deposited on glass) were measured by spectroscopic Ellipsometer (J. A. Woollam) using a black tape. Surface roughness (RMS) values of the Z-cpTi and cpTi were measured by surface profilometer (Model SE-2300, Kosaka). For the evaluation of surface roughness, 5 measurements over a scan length of 3–5 mm were performed and the average data was reported. Pencil hardness of the coated and uncoated cpTi surfaces was evaluated following ASTM D3363 specifications using a BYK Gardner pencil hardness tester. The cross-cut adhesive tape test of the nanocoating was carried out following ASTM D3359 specification to evaluate the adhesion of the coating with the Ti implant substrate. After the cross-cut adhesive tape test the coated surface was examined under an optical microscope using 100× magnifications. Field emission scanning electron microscopic (FESEM) analyses of the Z-cpTi and bare cpTi surfaces before and, after immersion in simulated body fluid (SBF) were done by Carl Zeiss, Germany, SUPRA-35VP instrument.

Evaluation of the apatite forming bioactivity

To investigate the bioactivity of the Z-cpTi samples, SBF was prepared according to the protocol of Kokubo et al. by dissolving of NaCl, KCl, MgCl₂·6H₂O, CaCl₂·2H₂O, Na₂SO₄, K₂HPO₄·3H₂O, NaHCO₃ and TRIS buffer [(CH₂OH)₃CNH₂] in ultrapure water.³⁰ To adjust the pH value (7.4 at 37 °C) 1 M HCl was used. The coated sample plates and uncoated reference plates (1.5 × 1 × 0.1 cm³) were immersed in SBF at 37 °C in polypropylene containers. The sample surface area (SA) to solution volume ratio (V) used for the static in vitro test was SA/V = 0.4 cm⁻¹.

MTT assay and SEM analysis using MG63 cell line

The MTT assay (Sigma, MO, USA) and SEM analyses were performed on cpTi and Z-cpTi substrates with pre-osteoblastic cell line MG63 (NCCS Cell Repository, Pune, India), derived from human osteosarcoma cells to investigate quantitatively the cytotoxicity as well as the cell proliferation, and visualize this primary cellular response, respectively.^31,32 The growth of viable MG63 cells (10³ cells were cultivated per substrate) attached on the sample surfaces after 3 and 5 d of incubation period was measured by this MTT assay and for comparison, cells cultured on blank wells were used as a control. Three samples for the each composition and two culture durations were used to evaluate cell proliferation. MTT solution [3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl tetrazolium] of 5 mg mL⁻¹ was prepared by dissolving MTT in PBS (phosphate buffer saline). Freshly prepared MTT solution was used for assay which was also protected from light. MTT solution was diluted (with DMEM culture media enriched with 10% fetal bovine so that the final concentration of MTT solution becomes 1 mg mL⁻¹) and added to each sample to form formazan salt, due to oxidation of tetrazolium by the action of mitochondrial dehydrogenase enzyme. After 4 h incubation at 37 °C inside the incubator, samples were transferred to a new 24-well plate and 1 mL of solubilisation solution (DMSO – dimethyl sulfoxide) was added to dissolve the formazan crystals. Then 100 μL of the resulting supernatant was transferred into each wells of a 96-well (at least three data points were obtained from each sample). The optical density of the solution in each well was measured at a wavelength of 570 nm using a micro plate reader (BioRad, USA). In the MTT assay study, cells cultured on tissue culture plate wells without any treatment/substrate and the same treated with 10% Triton-X 100 solutions were used as negative and positive control, respectively. OD of empty wells containing 100 μL DMSO was measured as blank measurements.

Parallely, on top of one of the samples (cpTi and Z-cpTi substrates), MG63 was grown for 3–5 days as mentioned. All the samples for SEM observation were first fixed with 2% paraformaldehyde in 0.1 M phosphate buffer saline overnight at 4 °C. The fixed samples were then dehydrated in series of ethanol (30, 50, 70, 95 and 100% three times), followed by gold sputter coating and observation under SEM (Phenom pro-X, Netherlands) for cell morphologies subsequently.

Results and discussion

Crystalline ZrO₂ nanocoatings on cpTi plates were prepared and quality of those coatings was measured by several standard tests. The exterior of bare cpTi and Z-cpTi samples showed surface roughness (root mean square; RMS) values of ≈470 and 590 nm, respectively indicating some enhancement of surface roughness of the implant due to the formation of crystalline ZrO₂ on the top. As known, the bare cpTi is greyish in colour (Fig. 1a). The single layer (1L) of ZrO₂ coating on this cpTi plate accomplished sufficient crystallinity after annealing at 500 °C for 2 h. However, the exposed and underlying Ti substrate turned bluish (Fig. 1b) due to the oxidation, thermal induced change or contaminations of Ti in presence of air which may diminish the original metallic strength, corrosion behaviour and other physical/biological properties of the implant. It should be noted here that we have also calcined these coated samples under N₂ (inert) atmosphere to obtain this cubic crystalline phase of ZrO₂ avoiding oxidative changes in Ti. However, even in nitrogen atmosphere the obvious thermal degradation, contamination and colour changes of Ti could not be restricted at such higher temperature. Further, heat-treatment in N₂ could cause C deposition. By lowering the temperature at 350 °C and increasing the annealing time up to 10 h, we tried to solve this problem. Interestingly, applying three consecutive layers (3L) of ZrO₂ with controlled thermal treatment, coating with an aesthetic look was achieved after annealing at lower temperature of 350 °C for 10 h without noticeable deterioration of the original cpTi substrate (Fig. 1c). Although a very faint coloration was observed in the uncoated part (Fig. 1c), the underneath cpTi surface remains unaffected. The nice appearance of the 3L coating with colour-hue is expected to be resulted due to the generation of fine micro/nanoscopic ZrO₂ crystals in the coating. Fig. 1d represents such fabricated final Z-cpTi plate heat-treated at 350 °C for 10 h in air. The lifting speed was optimized considering the most favourable thickness to achieve excellent crystallinity, adhesion and hardness of the final nanocoating. It is noteworthy here that application of withdrawal velocity higher than 13 cm min⁻¹ from the sol failed to generate desired crystallinity and adhesion of 3L-ZrO₂ nanocoating on Ti. The final Z-cpTi samples were subjected to the adhesion test (cross-cut adhesive tape test), and after the test the coated surfaces were examined under an optical microscope using 100× magnification. No peeling off of the nanocoating from the Ti substrate was observed. In addition, the Z-cpTi showed pencil hardness close to 9H which is very high compared to the bare cpTi surface (∼3H) and expected to be resulted mainly due to the elevated mechanical strength and toughness of the crystalline cubic ZrO₂ coating.^14,18 Such a high surface hardness of the coating can also be explained due to the sub-micron level unevenness of metallic Ti surface which helped in anchoring the ZrO₂ layer. It has been also established that the hcp (Ti) and fcc crystal phases (ZrO₂) are thermodynamically and structurally quite similar to each other which favours the growth of fcc on hcp phase.^33,34 Therefore, in this case it can be predicted that the hcp lattice of Ti metallic implant simultaneously facilitated the growth of fcc ZrO₂ crystalline phases all along on the Ti hcp facets resulting excellent atomic level adhesion and rigidity of the final nanocoating.

Fig. 1 Images of (a) uncoated cpTi; partially coated plates with (b) 1L-ZrO₂ heat-treated at 500 °C for 2 h and (c) 3L-ZrO₂ heat-treated at 350 °C for 10 h; (d) image of the fabricated final Z-cpTi (whole part is coated).

Structural and compositional analyses

The nature of crystalline phases affects the chemical inertness, physical property and biocompatibility of ZrO₂. Therefore, phase structure study of ZrO₂ nanocoating on cpTi plates were examined by high angle GIXRD measurements within the range from 25 to 80° 2θ. Low grazing incidence angle (0.3°) was maintained to restrict the penetration of X-ray within <400 nm which could cover well the coating layer, interface and a small part of the cpTi substrate.^35,36 The GIXRD pattern of 1L-ZrO₂ nanocoating on cpTi shows that after heat-treatment at 500 °C for 2 h in air, crystalline cubic ZrO₂ was formed but the peaks for underlying Ti reduced noticeably indicating the deterioration of original metallic properties of Ti (Fig. 2a). This damage in Ti due to this thermal treatment (500 °C/air) was also visually observed in the uncoated part (deep blue colour) of the image in Fig. 1b. To avoid oxidation of Ti, ZrO₂ coated cpTi substrates were heat-treated at lower temperature (350 °C/10 h). However, the desired crystallinity of ZrO₂ nanocoating was not observed after applying 1 (Fig. 2b) and 2L (Fig. 2c). Interestingly, in case of the 3L coated samples (Z-cpTi) the crystalline peaks of cubic ZrO₂ were generated after annealing at 350 °C in air for 10 h along with the retention of hexagonal Ti peaks of underlying substrate. Fig. 2d confirms the formation of cubic crystalline ZrO₂ in Z-cpTi with several prominent peaks at 30.33, 50.58 and 60.12 2θ corresponding to (111), (220) and (311) planes, respectively for cubic ZrO₂ phase (fcc).¹⁸ The generation of crystalline cubic ZrO₂ phase in Z-cpTi sample is well matched with JCPDS card no. 01-071-4810. Moreover, this XRD data (Fig. 2d) also reveals crystalline peaks for undamaged Ti metal at 38.39, 40.16, 52.98, 62.94, 70.63 and 76.19 2θ corresponding to (002), (101), (102), (110), (103) and (112) planes, respectively of hexagonal crystalline phase (hcp) of Ti (JCPDS card no. 01-089-3073). It can be anticipated that the layer-by-layer deposition restricted the overall thickness of the coating, and the hcp lattice structure of metallic Ti facilitated the growth of crystalline fcc ZrO₂ at such a low temperature. It should be noted here that applying single layer with a thickness equivalent to the overall thickness given in 3 layers resulted cracking of the coating. Therefore, in order to obtain homogeneous, compact and crack-free coating such layer-by-layer dip-coating approach followed by controlled thermal treatment was necessary. After investigating all the experimental parameters and subsequent characterizations it can be concluded that only the 3L-ZrO₂ coated Ti samples annealed at 350 °C for 10 h (Z-cpTi) was able to maintain the properties of crystalline cubic ZrO₂ layer with a nice look and strong adhesion on cpTi which can be utilized as a superior dental implant. It should be noted here that the high crystallinity of ZrO₂ could enhance the nucleation of the calcium phosphate on the implant upon soaking in body fluid.¹⁰

Fig. 2 GIXRD patterns of coated cpTi samples with (a) 1L-ZrO₂ heat-treated at 500 °C for 2 h in air; (b) 1L, (c) 2L, (d) 3L of ZrO₂ films heat-treated at 350 °C for 10 h in air. Fcc ZrO₂ and hcp Ti peak positions are also shown.

Since the Ti metal has a sub-micron level uneven surface, the thickness and refractive index measurement of the nanocoating on cpTi by ellipsometer was expected to be erroneous. Considering this fact, we have fabricated the 3L-ZrO₂ coating on smooth glass surface as a representative, and heat-treated maintaining the identical conditions. Though the dried coating was completely amorphous (Fig. S1a, ESI†), after annealing, similar type crystalline cubic ZrO₂ phase on glass (as shown Fig. S1b, ESI†) with moderate adhesion and surface hardness was also obtained. When the coating transformed into crystalline ZrO₂, the surface OH groups were decreased significantly and as a result, the glass-coating interface became weak. However, in case of Ti, this problem could not occur due to sub-micron level surface roughness and compatible hcp phase of Ti which were helping to anchor the coating. Nevertheless, as similar fcc phase of ZrO₂ was also formed on glass, we used this coating to evaluate the thickness and refractive index. Before heat-treatment, the as prepared 3L-ZrO₂ coating showed thickness and refractive index values of ∼156 nm and ∼1.62, respectively. However, after annealing (350 °C/10 h in air) the thickness and refractive index of the final ZrO₂ (measured by ellipsometer) nanocoating became ∼95 nm and ∼1.90, respectively.

FTIR studies of the dried and crystalline ZrO₂ nanocoating heat-treated at 350 °C was undertaken to understand the chemical bonding (Fig. 3). FTIR of 3L as-dried (70 °C) nanocoating showed bands at 1602, 1525 and 1384 cm⁻¹ which can be assigned due to the existence of acac chelate (Fig. 3a).^28,37 FTIR of the final heat-treated (350 °C/10 h) coating (Fig. 3b) showed complete removal of organics. The peaks at 650 and 430 cm⁻¹ (Fig. 3b) can be attributed to the vibration of Zr–OH and Zr–O–Zr linkages, respectively.^38,39 The band centered at 3431 cm⁻¹ is due to the OH stretching present on the surface of the coating material. These surface hydroxyl groups associated with ZrO₂ and few Zr–OH groups on Z-cpTi surface will possibly facilitate calcium phosphate nucleation in presence of body fluid.^10,11,21,40 During soaking in body fluid (pH ∼ 7.4), initially these surface OH groups after deprotonation absorb Ca²⁺ through electrostatic force, then HPO₄²⁻ available in fluid shifts towards the surface of the coated implant to attach with Ca²⁺ promoting the formation of apatite crystals.¹¹

Fig. 3 FTIR spectra of (a) dried and (b) heat-treated (350 °C/10 h) 3L-ZrO₂ nanocoating with peak assignments.

The surface morphology plays a crucial role in biological behaviour of implants. Therefore, the surfaces of Z-cpTi and cpTi samples were characterized by FESEM (Fig. 4). After depositing ZrO₂ nanocoating through layer-by-layer approach followed by thermal treatment at 350 °C for 10 h, a relatively homogeneous surface with micro/nano aggregates of crystalline material was observed (Fig. 4a and b). This nanostructured surface roughness is also expected to induce apatite nucleation and anchor cell proliferation on the implant during application in biological environment which is not possible by bare implant surface.⁴⁰ On the other hand, Fig. 4c shows metallic uneven surface of market available cpTi plate. The enhanced roughness in case of Z-cpTi than control cpTi is in agreement with the RMS value acquired from the surface profilometer.

Fig. 4 FESEM images of the (a and b) surface of the Z-cpTi with different magnifications showing the presence of aggregated ZrO₂ nanocrystals; (c) uneven surface of control cpTi metal surface with similar magnification of (a) is shown for comparison.

It is noteworthy here that the Z-cpTi coating surface is quite hard and the coating material could not be scratched off using steel knife. So, for TEM studies we used scratched off samples from the glass surface. The morphology of final crystalline ZrO₂ nanocoating material heat-treated at 350 °C for 10 h can be directly observed by TEM (Fig. 5a). SAED pattern with well-defined rings (inset in Fig. 5a) and the HR-TEM image (Fig. 5b) confirm the formation of crystalline cubic ZrO₂ phase in the matrix. The d-spacing (0.294 nm) calculated from HR-TEM fringes was well matched with the (111) plane of cubic ZrO₂ as also observed from XRD.¹⁸ The EDX pattern (Fig. 5c) acquired from the different portions of the TEM image reveals the presence of Zr and O elements in the nanocoating. Signals of Cu observed in the EDX are from Cu grid used for the TEM study. No crystalline phase was noticed in the as-dried (70 °C) nanocoating material before annealing (Fig. S2, ESI†) which is corroborated with the XRD pattern of the above dried nanocoating (Fig. S1a, ESI†).

Fig. 5 (a) TEM image showing crystalline ZrO₂ in nanocoating material heat-treated at 350 °C for 10 h. SAED pattern (inset in a) with the prominent rings and HR-TEM fringes (b) represent crystalline cubic ZrO₂ phase. (c) EDX spectrum shows the presence of Zr and O elements in the nanocoating.

In vitro apatite forming bioactivity of Z-cpTi

In vitro apatite forming ability is one of the typical characteristic of bioactive coatings which endorses bone-bonding ability of the implant.^21,22 The preparation process, crystalline phase pattern and surface microstructure influence the bioactivity of ZrO₂. To investigate the bioactivity of the newly designed Z-cpTi, samples were immersed in SBF and change in surface morphology was monitored (Fig. 6). The study was also performed with the uncoated cpTi plate as control maintaining the identical condition. From the FESEM image, island-like spherically shaped calcium phosphate deposition^10,11,23,26 was observed on Z-cpTi after 14 d of immersion (Fig. 6). Whereas, no precipitation was found on the control cpTi (inset in Fig. 6). Thus, FESEM study of Z-cpTi after immersion concludes that this nanostructured crystalline cubic ZrO₂ coating with sub-micron level surface roughness and surface hydroxyl groups (observed by FTIR) has tendency to facilitate calcium phosphate nucleation which could improve the acceptability of the coated dental implants.⁴⁰

Fig. 6 FESEM images of Z-cpTi and control cpTi (inset) plates after 14 days of immersion in SBF.

In addition, the XRD data (Fig. 7a) confirms that the precipitation formed on Z-cpTi after 14 d of immersion in SBF was carbonated hydroxyapatite (CO₃HA). The additional XRD peaks for crystalline CO₃HA are well matched with the JCPDS card no. 01-075-3727. The main peak observed at 2θ = 31.699° can be assigned as the (121) plane of CO₃HA phase.^21,22 No CO₃HA deposition was found in case of control cpTi plates (Fig. 7b). These observations are also well matched with the FESEM result of 14 d immersed samples (as shown in Fig. 6).

Fig. 7 XRD patterns of the Z-cpTi (a) and control cpTi (b) samples before and after 14 d of immersion in SBF.

Evaluation of the biocompatibility and cellular response

The human pre-osteoblastic MG63 cells used in this study for MTT assay are typically useful to assess the cytotoxicity of biomaterials which effects bone growth metabolism⁴¹ and cell–biomaterial interactions in bone tissue engineering.⁴² Bone tissue engineering materials should promote expression of osteoblastic phenotype through sequence of events, e.g., cell attachment and proliferation.⁴³ The biocompatibility and degrees of cell proliferation on the final Z-cpTi and control cpTi samples were quantitatively examined and compared after 3 and 5 d of culturing using an MTT assay (as shown in Fig. 8). From the calculated OD values, percentages cell proliferations were plotted against the days monitored and are given in Fig. 8. It was found that initially (3 d), the proliferation of cells on Z-cpTi was better than bare cpTi due to primary attachment of cells. This tendency was also continued as observed in 5 d. Interestingly, the cell proliferation was enhanced in case of the newly designed Z-cpTi compared to the control cpTi during 5 d of culturing suggesting its excellent biocompatibility, noncytotoxicity and cell attachment ability.

Fig. 8 Percentage of MG63 proliferation obtained from MTT assay on the Z-cpTi and reference cpTi samples incubated for 3 and 5 d.

Cell morphology acquired by SEM analysis (Fig. 9) also reveals similar trend as obtained from the MTT assay. SEM images exhibit better MG63 proliferation in case of Z-cpTi surface after 3 (Fig. 9a and b) and 5 d (Fig. 9c and d) of cell culturing than the control cpTi (Fig. 9e–h). Well-grown filopodia (microspikes) or cytoplasmic projections were seen and found to be more pronounced on newly designed Z-cpTi sample (Fig. 9b and d). Filopodia contain actin filaments cross-linked into bundles by actin-binding proteins. Several micro/nano aggregates present on the top of this surface played pivotal role for better anchorage of the filopodia. Osteoblastic proliferation may also depend on interaction (molecular and ionic) between biomaterial and its associated physiological environment that results release of specific ionic species, pH modifications and adsorption of biologically active molecules from such environment to cells' surface.⁴⁴ It is anticipated that due to the outstanding biocompatibility and chemical inertness of developed crystalline cubic ZrO₂ on the surface, such improvement in cell attachment ability of the coated dental implant was gained. In addition, we have obtained higher surface roughness of Z-cpTi plates compared to the bare cpTi which is also one of the major factors influencing the cell adhesion and proliferation of coated implant's surface.⁴⁵

Fig. 9 The SEM images with different magnifications reveal cell growth and morphology on final Z-cpTi after 3 (a and b) and 5 d (c and d) of cell culturing. The surfaces of cell cultured cpTi implants after 3 (e and f) and 5 d (g and h) are also shown for comparison.

Schematic in Fig. 10, demonstrates the entire experimental procedure of Z-cpTi development and its superiority as advanced dental implant. The fine crystalline Z-cpTi was produced applying a controlled layer-by-layer ZrO₂ nanocoating deposition on cpTi followed by drying and heat-treatment at low temperature. The cpTi without any surface modification showed no apatite forming bioactivity and failed to achieve enhanced biocompatibility. Whereas, the nanostructured Z-cpTi surface induced the nucleation of apatite layer in SBF by the help of bioactive crystalline cubic ZrO₂. Furthermore, generated surface and excellent biocompatibility of cubic ZrO₂ on Z-cpTi facilitated attachment, and growth of the cells in a quite large number compared to the market available cpTi implant.

Fig. 10 Schematic represents the fabrication procedure of Z-cpTi with enhanced biocompatibility and exclusive apatite forming ability compared to the bare cpTi dental implant.

Conclusions

Crystalline cubic ZrO₂ nanocoating with excellent surface hardness and adhesion was fabricated in situ on commercially pure titanium metal (market available dental implant, grade 4) for application as improved tooth replacement implants. The crystallinity of the nanocoating was achieved at lower annealing temperature following a controlled layer-by-layer dip-coating technique and thermal treatment. This novel and attractive synthetic approach helped to improve the chemical, physical as well as biological property of the implant by depositing crystalline ZrO₂ nanocoating on the surface without any noticeable damage of the original properties of metallic Ti. The newly designed chemically inert nanostructured cubic ZrO₂ coating exhibited appreciable bioactivity in terms of apatite depositing ability, excellent biocompatibility and tissue attachment property which could enhance the osseointegration as well as reduce inflammatory reaction of the metallic oral implants. Moreover, the corrosion resistant property, high surface hardness and aesthetic appearance of Z-cpTi can establish it as pivotal towards the fabrication of a long-lasting advanced dental implant.

Acknowledgements

The research was financially supported by the Department of Science and Technology, DST, India (Project No. INT/FINLAND/P-11). ID and SC thank CSIR and UGC for awarding fellowships.

References

1. R. Family, M. Solati-Hashjin, S. N. Nik and A. Nemati, Caspian J. Intern. Med., 2012, 3, 460–465.
2. Z. Ozkurt and E. Kazazoglu, J. Oral Implantol., 2011, 37, 367–376.
3. R. B. Osman and M. V. Swain, Materials, 2015, 8, 932–958.
4. A. Apratim, P. Eachempati, K. K. K. Salian, V. Singh, S. Chhabra and S. Shah, J. Int. Soc. Prev. Community Dent., 2015, 5, 147–156.
5. R. Castellani, A. D. Ruijter, H. Renggli and J. Jansen, Clin. Oral Implants Res., 1999, 10, 369–378.
6. G. Thrivikraman, G. Madras and B. Basu, RSC Adv., 2014, 4, 12763–12781.
7. J. M. Gomez-Vega, E. Saiz, A. P. Tomsia, G. W. Marshall and S. J. Marshall, Biomaterials, 2000, 21, 105–111.
8. M. F. Morks, Mater. Lett., 2010, 64, 1968–1971.
9. W. K. Yeung, I. V. Sukhorukova, D. V. Shtansky, E. A. Levashov, I. Y. Zhitnyak, N. A. Gloushankova, P. V. Kiryukhantsev-Korneev, M. I. Petrzhik, A. Matthews and A. Yerokhin, RSC Adv., 2016, 6, 12688–12698.
10. A. Thaveedeetrakul, V. Boonamnuayvitaya and N. Witit-anun, Adv. Mater. Phys. Chem., 2012, 2, 45–48.
11. L. Guo, J. Zhao, X. Wang, X. Wang, X. Xu, H. Liu and Y. Li, Int. J. Appl. Ceram. Technol., 2009, 6, 636–641.
12. T. R. Ramesh, M. Gangaiah, P. V. Harish, U. Krishnakumar and B. Nandakishore, Trends Biomater. Artif. Organs, 2012, 26, 154–160.
13. A. A. Madfa, F. A. Al-Sanabani, N. H. Al-Qudami, J. S. Al-Sanabani and A. G. Amran, Open Biomater. J., 2014, 5, 1–9.
14. K. Harada, A. Shinya, D. Yokoyama and A. Shinya, J. Prosthodont. Res., 2013, 57, 82–87.
15. B. Masheder, C. Urata and A. Hozumi, ACS Appl. Mater. Interfaces, 2013, 5, 7899–7905.
16. C. Hubsch, P. Dellinger, H. J. Maier, F. Stemme, M. Bruns, M. Stiesch and L. Borchers, Acta Biomater., 2015, 11, 488–493.
17. P. Palmero, M. Fornabaio, L. Montanaro, H. Reveron, C. Esnouf and J. Chevalier, Biomaterials, 2015, 50, 38–46.
18. D. Tan, G. Lin, Y. Liu, Y. Teng, Y. Zhuang, B. Zhu, Q. Zhao and J. Qiu, J. Nanopart. Res., 2011, 13, 1183–1190.
19. H. Ehrlich, P. Simon, M. Motylenko, M. Wysokowski, V. V. Bazhenov, R. Galli, A. L. Stelling, D. Stawski, M. Ilan, H. Stocker, B. Abendroth, R. Born, T. Jesionowski, K. J. Kurzydlowski and D. C. Meyer, J. Mater. Chem. B, 2013, 1, 5092–5099.
20. G. Sponchia, E. Ambrosi, F. Rizzolio, M. Hadla, A. D. Tedesco, C. R. Spena, G. Toffoli, P. Riello and A. Benedetti, J. Mater. Chem. B, 2015, 3, 7300–7306.
21. I. Das, S. K. Medda, G. De, S. Fagerlund, L. Hupa, M. A. Puska and P. K. Vallittu, J. Am. Ceram. Soc., 2015, 98, 2428–2437.
22. I. Das, G. De, L. Hupa and P. K. Vallittu, Mater. Sci. Eng., C, 2016, 62, 206–214.
23. M. Uchida, H.-M. Kim, F. Miyaji, T. Kokubo and T. Nakamura, Biomaterials, 2002, 23, 313–317.
24. H.-L. Huang, Y.-Y. Chang, J.-C. Weng, Y.-C. Chen, C.-H. Lai and T.-M. Shieh, Thin Solid Films, 2013, 528, 151–156.
25. V. Sollazzo, F. Pezzetti, A. Scarano, A. Piattelli, C. A. Bignozzi, L. Massari, G. Brunelli and F. Carinci, Dent. Mater., 2008, 24, 357–361.
26. G. Wang and H. Zreiqat, Materials, 2010, 3, 3994–4050.
27. M. Berni, N. Lopomo, G. Marchiori, A. Gambardella, M. Boi, M. Bianchi, A. Visani, P. Pavan, A. Russo and M. Marcacci, Mater. Sci. Eng., C, 2016, 62, 643–655.
28. G. De, A. Chatterjee and D. Ganguli, J. Mater. Sci. Lett., 1990, 9, 845–846.
29. G. I. Spijksma, H. M. Bouwmeester, D. H. A. Blank and V. G. Kessler, Chem. Commun., 2004, 1874–1875.
30. T. Kokubo and H. Takadama, Biomaterials, 2006, 27, 2907–2915.
31. S. Nagarajan, M. Mohana, P. Sudhagar, V. Raman, T. Nishimura, S. Kim, Y. S. Kang and N. Rajendran, ACS Appl. Mater. Interfaces, 2012, 4, 5134–5141.
32. X. Wang, Y. Li, P. D. Hodgson and C. Wen, Tissue Eng., Part A, 2010, 16, 309–316.
33. J. P. Hoogenboom, A. K. V. Langen-Suurling, J. Romijn and A. V. Blaaderen, Phys. Rev. Lett., 2003, 90, 138301.
34. J. Russo and H. Tanaka, Soft Matter, 2012, 8, 4206–4215.
35. Y. Cai, Y. Pan, J. Xue and G. Su, Appl. Surf. Sci., 2009, 255, 4066–4073.
36. S. Sembiring, B. O'Connor, D. Li, A. V. Riessen, C. Buckley and I. Low, Adv. X-Ray Anal., 2000, 43, 319–325.
37. I. Das and G. De, Sci. Rep., 2015, 5, 18503.
38. C. J. Fu, Z. W. Zhan, M. Yu, S. M. Li, J. H. Liu and L. Dong, Int. J. Electrochem. Sci., 2014, 9, 2603–2619.
39. M. R. Elvira, M. A. Mazo, A. Tamayo, F. Rubio, J. Rubio and J. L. Oteo, J. Chem. Chem. Eng., 2013, 7, 120–131.
40. G. Wang, F. Meng, C. Ding, P. K. Chu and X. Liu, Acta Biomater., 2010, 6, 990–1000.
41. Q. Wang, S. Zhong, J. Ouyang, L. Jiang, Z. Zhang, Y. Xie and S. Luo, Clin. Orthop. Relat. Res., 1998, 348, 259–268.
42. C. Wirth, B. Grosgogeat, C. Lagneau, N. Jaffrezic-Renault and L. Ponsonnet, Mater. Sci. Eng., C, 2008, 28, 990–1001.
43. D. D. Deligianni, N. D. Katsala, P. G. Koutsoukos and Y. F. Missirlis, Biomaterials, 2000, 22, 87–96.
44. M. P. Ferraz, J. C. Knowles, I. Olsen, F. J. Monteiro and J. D. Santos, Biomaterials, 2000, 21, 813–820.
45. M. Gahlert, T. Gudehus, S. Eichhorn, E. Steinhauser, H. Kniha and W. Erhardt, Clin. Oral Implants Res., 2007, 18, 662–668.

---

Footnote

† Electronic supplementary information (ESI) available: GIXRD patterns of dried and heat-treated 3L-ZrO₂ coating on glass (Fig. S1), TEM with SAED pattern of the dried ZrO₂ coating material (Fig. S2). See DOI: 10.1039/c6ra10661g

---