Retracted Article: Smart rose flower like bioceramic/metal oxide dual layer coating with enhanced anti-bacterial, anti-cancer, anti-corrosive and biocompatible properties for improved orthopedic applications

Abstract

Metallic implants suffer from numerous problems such as stress shielding, poor prolonged osseointegration and corrosion under in vivo environments. Such problems are often faced by bone cancer patients as they receive orthopedic implants after cancerous bone resection. Unfortunately, there are no orthopedic materials developed so far that simultaneously increase healthy bone growth as takes place in traditional orthopedic implant applications, while inhibiting cancerous bone growth. Based on these issues, the long-term objective of this study was to introduce a new implant material in an integrated way. Hence we have fabricated a selenium (Se), and manganese (Mn) substituted flower like hydroxyapatite (HAP) coating on zirconium oxide (ZrO₂) coated AZ91 magnesium alloy. The flower like Se,Mn-HAP/ZrO₂ dual layer coating on the AZ91 magnesium alloy was characterized by Fourier transform infrared spectroscopy (FTIR), X-ray diffraction (XRD), high-resolution scanning electron microscopy (HRSEM) and energy dispersive X-ray (EDX) analysis. Also, the mechanical properties of the dual layer coating were evaluated using adhesion and Vickers micro-hardness tests. The effect of the dual layer coating on the corrosion behavior of the AZ91 magnesium alloy was also investigated in simulated body fluid (SBF) using electrochemical studies. The cell–material interaction of the dual layer coating was observed in vitro with human osteosarcoma MG63 cells for cell proliferation at 1, 4 and 7 days of incubation and in vivo in Wistar rats for 14 and 28 days of implantation. From the results it was found that the dual layer coated AZ91 Mg alloy has an optimal structure and morphology, as indicated by SEM, showing a desired surface for the osteoblast adhesion, viability and proliferation. The new ceramic coating has induced an increased adhesion strength and microhardness, improved corrosion resistance, enhanced osteoblast proliferation and inhibited the growth of cancerous cells. Therefore, based on these results, we propose a new dual layer coated AZ91 Mg alloy which satisfies the requirements in bone cancer treatment and signifies progress in the field of implant materials.

---

1. Introduction

Bones are a common place for cancer cells to settle and start growing. Cancer cells that break off from a primary tumor and enter the bloodstream or lymph vessels can reach nearly all tissues of the body.¹ Tumors that result from these cells entering the bones are called bone metastases. Further, bone cancer describes cancer cells that originate in the bone itself. Bone cancer can cause bone pain when the cancer disrupts or destroys the bone's normal structure.² Such problems with promoting new bone growth next to implant surfaces are made more complex for patients with bone cancer (both primary bone cancer and metastasized bone cancers). For these reasons, the main goal of this study was to introduce a new bone replacing metallic biomaterial.^3,4 The use of metallic implants for a whole range of medical implants is justified by their superior mechanical properties, biocompatibility, etc.⁵ Among the available metallic implant materials, magnesium and its alloys have been broadly considered for bone replacing implants, due to several properties such as lightweight, high strength to weight ratio, low Young's modulus equivalent to that of human bone. That makes them promising candidates for biomedical applications.^6,7 These materials exhibit biocompatibility and appropriate mechanical properties for use as an orthopedic implant.⁸ Further, Mg and Mg alloys can corrode easily under physiological conditions such as in human body fluid or blood plasma solution which opens the door for second surgical intervention to remove the implant. Moreover, the fast or uncontrolled corrosion is usually associated with strong hydrogen ion release and severe pH changes, which pave the way for various undesirable biological reactions. In parallel to this, a loss of mechanical stability before the complete bone healing is observed. So it is necessary to improve the corrosion resistance of Mg based materials for bone implant application. Many researchers have pointed out that the surface modification with an appropriate coating is an effective approach, which could improve the corrosion resistance and surface biocompatibility of Mg based implants.^9–11

Furthermore for patients suffering from bone cancer, the cancerous bone was commonly replaced with an orthopedic implant in clinic. However, there is still a risk of cancer recurrence after the implantation due to the latency of cancer cells. Meanwhile, bacteria-induced inflammatory reaction is also one of the main factors that lead to implant loosening after implantation. Therefore, it is urgent to develop the functional bone implants with anticancer and antibacterial properties.¹²

Hydroxyapatite (HAP) bioceramics is one of basic biomaterials that is used in bone implant surgery.¹³ Due to high biocompatibility, bioactivity and very good adaptation under in vivo conditions, they are widely used in treating bone defects in the field of orthopedics.¹⁴ Further, the HAP based bioceramic material might be improved by the substitution of various bioactive ions to enhance the biological properties of the implants.^15,16 In particular, the favorable effect of bioactive ions such as strontium (Sr²⁺), magnesium (Mg²⁺), manganese (Mn²⁺), silver (Ag²⁺), zinc (Zn²⁺), silicon (Si²⁺), carbonate (CO₃²⁻), etc., in HAP based coatings on metallic substrates has been reported.^17,18 Each mineral ion has unique properties like enhancing the physicochemical, mechanical and biological properties of apatites and based on the need, the specific elements can be substituted into the apatite. Since our present study is focused on the development of a bioceramic material that is to be used for the bone cancer treatment, minerals with better anti-cancer property and bone bonding ability is substituted into the HAP lattice. Selenium (Se) that has chemopreventive properties is one such element which can satisfy the need of our material.^19,20

Until recently, selenium was considered as a highly toxic element, and hence harmful to human health.²¹ Se is an essential and unique trace element that plays a crucial role in health and disease. Meanwhile, Se has also been reported as an efficient anticancer agent.^22–24 Se inhibited the growth of cancerous osteoblasts and promoted the proliferation of healthy osteoblasts.^25,26 Numerous studies indicate that a selenium deficiency may cause an inhibition of bone growth in rats as well as a severe reduction of bone strength.²⁷ Apart from the therapeutic properties mentioned, Se also enhances the corrosion resistance of Cr, Pt and Mg alloys. Moreover, it is reported that Se substituted HAP is potentially a promising bone graft material.²⁸ In order to enhance the bone bonding ability and to promote osteoblast differentiation, a second substituent (Mn) is introduced into the coating.²⁹ Manganese is a very essential element for the growth and development of bones.³⁰ Since, the radius of the manganese ions (0.99 Å) is very close to that of calcium (0.90 Å), Mn ions easily enter osteoblasts through calcium ion channels. Manganese influences bone metabolism via the regulation of osteoblast differentiation and bone resorption.^31,32 In addition to the above property, Mn is essential for normal skeletal development, a well-functioning immune system, efficient hematopoiesis, healthy cartilage and bone tissue synthesis.³³ Thus, manganese has been substituted in HAP coatings to improve the biocompatibility and osteoblast differentiation. More significantly, the presence of Mn in calcium phosphate can also improve the production of osteocalcin in osteoblasts, even more effectively than the other minerals like strontium, magnesium, etc.³⁴ Although, the minerals substituted biomaterial is capable for improving various biological property, they could not maintain long-term stability, and may delaminate from the surface of the metallic implants in certain situations due to the lack of bonding strength, which in turn may lead to clinical complications and implant failure.³⁵

Hence, it is necessary to improve the bonding between the substrate and the coating. Moreover, Mg alloy possesses high corrosion rate in physiological medium and hence it must be protected from corrosion for its long-term application in the field of orthopedics.³⁶ To improve the bonding ability and to enhance the corrosion resistance, several techniques like surface modifications or surface treatments have been adopted. In the present work we have developed a zirconium oxide coating (ZrO₂) which has several advantages over other materials.^37,38 ZrO₂ has been used as a reinforcement phase due to its special mechanical properties, high corrosion resistance, physiological stability, biocompatibility and bioactivity.^39–42 Since biomaterials are permanent implant-materials, ZrO₂ can provide a much superior long-term corrosion protection to AZ91 Mg alloy. ZrO₂ has excellent technological and characteristics properties namely chemical and thermal stability, wear resistance, high strength and fracture toughness, and bioinertness.^43–48 Hence, the main aim of our work is the development of Se,Mn-HAP coating on ZrO₂ coated AZ91 Mg alloy. The HAP/ZrO₂ composite material may be an ideal biocomposite for orthopedic implantation because of its improved strength, hardness, and toughness where compared to the pure HAP.²⁵ In the present work, the first layer, i.e., ZrO₂ coated AZ91 Mg alloy improves the corrosion resistance and the second layer, i.e., Mn,Se-HAP coating which is flower-like in nature, certainly paves the way for the development of bone tissues through pores in between them.

A number of coating technologies are available for the development of bioceramic coatings and among them the electrochemical deposition has unique advantages due to its relatively low deposition temperature, process simplicity, capability of forming a uniform coating on porous substrates (or) complex shapes of subsets and the availability and low cost of equipments.^49–51 To the best of our knowledge, there are no reports on the electrodeposition of rose flower like Se,Mn-HAP coating on ZrO₂ coated AZ91 Mg alloy. Hence we propose that the Se,Mn-HAP coating on ZrO₂ coated AZ91 Mg alloy could improve the corrosion resistance, proliferation of healthy osteoblasts and inhibit the growth behaviors of cancer cells and bacteria cells.

2. Materials and methods

2.1 Mg alloy surface preparation

The elemental compositions (wt%) of the AZ91 Mg alloy used in this study are as follows: 0.59% Zn, 0.17% Mn, 8.63% Al, <0.05% Cu, <0.05% Fe and balance Mg. Samples were cut into 10 × 10 × 5 mm³ squares and then polished mechanically with various grades of SiC abrasive papers (180, 280, 360, 600 up to 1000). After polishing, all the samples were ultrasonically cleaned in 100% acetone for 10 min, in order to remove any surface residues. Finally the samples were rinsed in deionized (DI) water and then dried immediately in flowing air. The as cleaned AZ91 Mg alloys were placed under UV radiation for 1 h on each side for sterilization and then used for further studies.

2.2 Preparation of electrolyte

Analytical grade calcium nitrate dihydrate (Ca(NO₃)₂·2H₂O), sodium selenite Na₂SeO₃·5H₂O and manganese nitrate (Mn(NO₃)₂·6H₂O) were dissolved in deionized water separately and the solutions were mixed in the ratio of 8 : 1 : 1, respectively. The diammonium hydrogen phosphate ((NH₄)₂HPO₄) was dissolved in deionized water and both the solutions were mixed to produce the target (Ca + Mn)/(P + Se) ratio of 1.67 which is close to that of hydroxyapatite (HAP). The electrolyte was prepared in the N₂ gas atmosphere and the pH of the electrolyte was adjusted to 4.5 using dilute NH₄OH (or) HCl. The electrolyte was magnetically stirred for 2 h at a speed of 180 rpm to maintain the uniform concentration. In addition, pure HAP and Mn-HAP electrolyte was also prepared using Ca(NO₃)₂·2H₂O, Mn(NO₃)₂·6H₂O and (NH₄)₂HPO₄ by adopting the same procedure.

For the preparation of electrolyte for ZrO₂ coating, an appropriate amount of zirconium nitrate (Zr(NO₃)₂) was dissolved into 100 mL of distilled water in an airtight container and stirred continuously for 30 min at room temperature (28 ± 1 °C) to obtain an uniform and homogeneous transparent solution as precursor. All the chemicals were of analytical grade and used without further purification. Deionized water was employed as the solvent throughout the experiments.

2.3 Electrochemical deposition

2.3.1 Electrodeposition of ZrO₂ on AZ91 Mg alloy. The electrodeposition process of ZrO₂ on AZ91 Mg alloy was carried out using a regular three electrode cell arrangement by galvanostatic method using an electrochemical workstation (CHI 760C (CH Instruments, USA)) in which the platinum electrode served as the counter electrode, AZ91 Mg alloy and saturated calomel electrode (SCE) as the working and reference electrodes, respectively. The electrodeposition was carried out in a concentrated solution of 0.01 M Zr(NO₃)₂ at three different deposition time of 5, 10 and 15 min, respectively. After the electrodeposition process, the ZrO₂ coated AZ91 Mg alloy surface was washed with deionized water to remove residual electrolyte, then naturally dried for 24 h and were stored in a desiccators at room temperature.

2.3.2 Electrodeposition of HAP, Mn-HAP, Se,Mn-HAP on ZrO₂ coated AZ91 Mg alloy. The electrodeposition of HAP, Mn-HAP, Se,Mn-HAP, respectively was performed in a common three electrode configuration as described in previous section. The deposition was performed on AZ91 Mg alloy at the best current density of 9 mA cm⁻² which was optimized from the author's previous work.⁵² After the electrodeposition, coated specimens were removed from the electrolyte and were gently rinsed with deionized water, and then dried for 24 h. Finally, all the samples were sterilized in an autoclave and stored in desiccators at room temperature.

2.4 Characterization of the as developed coatings on AZ91 Mg alloy

2.4.1 Phase structure and morphological characterization. The functional characteristics of the as developed coatings were analyzed using Fourier transform-infrared spectroscopy (FT-IR, Impact 400 D Nicolet Spectrometer) to identify and confirm the functional groups. The spectra were recorded over the frequency range from 4000 cm⁻¹ to 400 cm⁻¹ with a number of 32 scans and spectral resolution of 4 cm⁻¹.

The phase composition of the coatings was analyzed by X-ray diffractometry. The crystalline structures and phase compositions of these coatings were characterized by X-ray diffraction (XRD, Seifert, X-ray diffractometer Siemens D500 Spectrometer) in the range between 20° ≤ 2θ ≤ 60° with CuKα radiation generated at 40 kV and 30 mA with a step size of 0.02° at a scanning rate of 10 min⁻¹.

The surface morphology of the experimental samples were observed using a field-emission scanning electron microscope Hitachi (FE-SEM S4800). Before the analysis, all the coated samples were sputtered with Au. The elemental analysis and elemental mapping were performed on dual coated AZ91 Mg alloy using FE-SEM to examine the distribution of various elements in the coatings.

2.5 Mechanical characterization

The adhesion strength of the coating is of utmost importance for the implant to function properly in load bearing applications. Adhesion strength of the dual layer coatings AZ91 Mg alloy specimens were examined by pull-out test according to the American Society for Testing Materials (ASTM) international standard F1044 with at least five measurements for each experiment. Prior to the adhesion test all the coated samples were cured in an oven at 150 °C for 1 h and the fixtures were subjected to pull-out test using a universal testing machine (Model 5569, Instron) at a crosshead speed of 1 mm min⁻¹.

Vickers micro-hardness tests were performed on the dual layer coated AZ91 Mg alloy samples using Akashi AAV-500 series hardness tester the universal test machine for the evaluation of hardness of the coatings. The force of the applied load used was 490.3 mN for a dwell time of 30 s. An average of each sample was subjected to at least five measurements under the hardness testing.

2.6 Electrochemical investigation

The corrosion resistance of the dual layer coated and uncoated AZ91 Mg alloy was evaluated by potentiodynamic polarization and electrochemical impedance spectroscopic (EIS) studies in the SBF solution which was prepared in accordance with Kokubo's recipe, in which the ion concentrations were similar to that of the human blood plasma⁵³ (Table 1). The electrochemical experiments were carried out using a standard three-electrode cell assembly, in which the platinum electrode, saturated calomel electrodes and AZ91 Mg alloy were used as counter, reference and working electrode, respectively. The studies were performed using CHI 760C Electrochemical Workstation. The potential of working electrode was measured against saturated calomel electrode (SCE). The 1 cm² surface area of the sample was kept in contact with the test solution. The potentiodynamic polarization measurements were performed from −400 mV to −2000 mV vs. SCE with a scan rate of 1 mV s⁻¹. The EIS measurements were performed with frequency ranging from 10⁻¹ Hz to 10⁵ Hz and with perturbation amplitude of 5 mV. The obtained data were recorded using the internally available software and each experiment was repeated three times to check the reproducibility.

Table 1 Ionic composition of human blood plasma and SBF⁵³

Ions	Na^+	K^+	Ca^2+	Mg^2+	Cl^−	HCO3^−	HPO4^2−	SO4^2−
Human blood plasma (mM)	142	5	2.5	1.5	103	27	1	0.5
SBF (mM)	142	5	2.5	1.5	147.8	4.2	1	0.5

---

2.7 Antibacterial activity

The in vitro antibacterial activity of the samples was tested against two bacterial strains E. coli and S. aureus by the agar disc diffusion method. The inocula of the selected bacterial species were prepared from the fresh overnight broth cultures (Tripton soy broth with 0.6% yeast extract-Torlak, Serbia) that were incubated at 37 °C with constant stirring and were then used for the diffusion studies. The diffusion technique was carried out by pouring agar into Petri dishes to form 4 mm thick layers and adding dense inocula of the test organisms of two bacterial strains in order to obtain better growth. Petri plates were left for 10 min in the laminar air flow and after that, discs (6 mm) were prepared from Whatman no. 3 filter paper, immersed into different volumes of (25, 50, 75, 100, 125 (μg mL⁻¹)) dual layer coated samples, placed at equal distances and then incubated at 37 °C for overnight in a bacteriological incubator. The results were obtained by measuring the width of zone of inhibition (mm) around the disc which was produced by the coating samples against the two bacterial strains.

2.8 Evaluation of viability of human fibroblast cells on dual layer coated Mg alloy

The cell viability of the dual layer coated AZ91 Mg alloy, was assessed using MTT assay (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide). The MTT assay test was performed in order to measure the cell viability. For this purpose, human fibroblast cells (L929, ATCC CRL-1427TM) supplied by National Centre for Cell Sciences (NCCS), Pune, India, were cultured in standard culture medium, Dulbecco's Modified Eagle Medium (DMEM, GIBCO), which consisted of a minimal essential medium, supplemented with 10% fetal bovine serum (FBS) and 1% non-essential amino acids (GIBCO). Cell suspensions were distributed in a 24-well plate at a density of 5 × 10³ cells per well in medium to determine the cell proliferation and cytotoxicity, as a function of incubation time of 1, 4 and 7 days. After the cells were seeded and incubated for 1, 4, and 7 days the samples were taken from the 24-well plates and rinsed with phosphate buffer saline thrice. The medium was renewed for every 2 days and the cultures were maintained in a humidified atmosphere with 5% CO₂ and 95% humidified air, at 37 °C in a CO₂ incubator. The solution was then removed, dimethyl sulfoxide was added to it, and the plate was shaken for 15 min before measuring absorbance at 570 nm on an ELISA microplate reader and then % of cell viability was calculated with respect to control.

2.9 Evaluation of cytotoxicity of bone cancer Saos-2 cells on dual layer coated Mg alloy

The human large cell bone cancer Saos-2 cells were purchased from the National Centre for Cell Sciences (Pune, India). Saos-2 cells were maintained in Dulbecco's Modified Eagles (DMEM) with 10% fetal bovine serum and were cultured in RPMI 1640 (Hi-Media, Inc.) supplemented with 10% fetal bovine serum and 1% penicillin and streptomycin. The cell lines were maintained at 37 °C with 5% CO₂ in a humidified incubator. Cell viability was determined by 3-(4,5-dimethylthiazol-2yl)-2,5-diphenyltetrazolium bromide (MTT) assay. Briefly, a density of 5 × 10³ cells per well in medium were seeded into a 24 well plate and allowed to attach overnight. Cells were treated with the dual layer coated Mg alloy samples and left contact for 1 day, 4 days and 7 days at 37 °C. Cell viability was calculated as follows.

2.10 Cell adhesion test

Human osteoblast like MG63 cells were grown in Dulbecco's Modified Eagle Medium (DMEM, GIBCO) added with 10% fetal bovine serum (FBS) and 1% penicillin streptomycin cocktail. The Mn,Se-HAP/ZrO₂ dual layer coated Mg alloy sample was used for cell adhesion experiment. The coated sample was sterilized in an autoclave at 121 °C for 25 min and the cells were seeded on the samples with the density of 5 × 10⁵ mL⁻¹ and then culture plates were placed in an incubator under standard cell culture conditions (37 °C, 5% CO₂ and 95% humidified atmosphere conditions air). After the stipulated time period of 36–48 h, the coated samples were rinsed with phosphate buffer saline (PBS) of pH 7.4. The MG63 cells attached to the coated sample were fixed with 2% glutaraldehyde for 1 h and cleaned twice with PBS. The dual layer coated sample was dehydrated with different grades of ethyl alcohol (30, 40, 50, 60, 70, 80, 90, and 100%) for 15 min and sputter covered with gold film and then morphology of the cells were examined under HRSEM.

2.11 Animal study and surgical procedure

The care for the animals and its surgical procedure were conducted as being compliant with the guidelines of the Institutional Animal Ethics Committee (IAEC) at Kovai Medical Center Research and Educational Trust, Coimbatore, India (KMCRET/Ph.D/3./2013–2014). A total of 10 male Wistar rats (5 animals per group) group I uncoated and group II dual layer coated. Weighing 200–250 g were housed under standard conditions at a controlled temperature (20 °C) and a light/dark cycle (12/12 h). Rats were individually anesthetized via intra peritoneal injections of ketamine (20 mg kg⁻¹) and xylazine (2 mg kg⁻¹) and inhaled a mixture of 20% v/v isoflurane and propylene glycol. The surgical site was first shaved, scrubbed with iodine and using blunt dissection, the muscles were separated over the femur to expose the periosteum. The implantation was conducted under aseptic condition. An incision was made at the femur bone by making a 2 mm hole using the drilling machine with sufficient irrigation with saline solution throughout the drilling process to minimize the temperature rise in the bone. The incisions were finally closed using absorbable sutures and was then allowed to recover from anesthesia in a warm environment while being observed. The animals received antibiotics (penicillin) for three postoperative days. The rats were kept on rearing for 14 and 28 days and the animals were monitored twice in a day, especially during the first week after surgery. Rats were sacrificed after 14 days and 28 days of implantation and specimens were harvested for the histological evaluation.

For histological observations, the bones with the implants were sectioned to a thickness of 1–2 mm, with a low speed diamond saw. Then the samples were fixed in neutral formalin solution of 20%, then decalcified inside acetic acid solution of 10% for 4 days and embedded in paraffin. Ultrathin sections were cut at 70 nm, but then bars on the rats' bones were removed before. Sections were stained with Mallory and hematoxylin eosin staring solution and histologically analyzed by light microscopy.

3. Results and discussion

3.1 Surface characterization of the dual layer (Se,Mn-HAP/ZrO₂) coatings

3.1.1 FT-IR analysis. The FT-IR analysis of HAP, Se,Mn-HAP, ZrO₂ and Se,Mn-HAP/ZrO₂ dual layer coated AZ91 Mg alloy was carried out and the corresponding spectra are shown in Fig. 1(a–d). The FT-IR spectrum of the HAP coating is shown in Fig. 1(a) which shows the characteristic peaks corresponding to phosphate (PO₄³⁻) groups (ν₁-957 cm⁻¹, ν₃-1030 and 1091 cm⁻¹, ν₂-479 cm⁻¹, ν₄-561 and 607 cm⁻¹) and hydroxyl (OH⁻) groups (stretching vibration at 3570 cm⁻¹ and libration mode at 631 cm⁻¹), respectively. Additionally, the peaks due to the stretching and bending modes of adsorbed water (H₂O) molecules can be seen at 3423 cm⁻¹ and 1629 cm⁻¹. The peak assignments are in accordance with the literature data. The spectrum of Se,Mn-HAP coating (Fig. 1(b)) shows a similar structure to that of the HAP coating (Fig. 1(a)), but with a slight change in the wave numbers of peaks. The peaks located at 3485 cm⁻¹ and 1638 cm⁻¹ are owing to the stretching and bending modes of adsorbed H₂O molecules, while the OH⁻ stretching and bending bands for Se,Mn-HAP are shifted to 3630 cm⁻¹ and 626 cm⁻¹, respectively. Whereas, the absorption peak due to PO₄³⁻ groups for Se,Mn-HAP can be clearly observed at 1032 and 1098 cm⁻¹ (ν₃), 559 cm⁻¹ and 598 cm⁻¹ (ν₄), 474 cm⁻¹ (ν₂) as well as 947 cm⁻¹ (ν₁), respectively. The peak at 773 cm⁻¹ is due to the O–Se–O bending vibration.

Fig. 1 FT-IR spectra of (a) HAP (b) Se,Mn-HAP (c) ZrO₂ and (d) Se,Mn-HAP/ZrO₂ dual layer coated AZ91 Mg alloy.

The FT-IR spectrum for the ZrO₂ coated AZ91 Mg alloy (Fig. 1(c)) exhibits the corresponding characteristic peak at 418 cm⁻¹ which gives an indication of presence Zr–O stretching bond. In addition to that, the peaks observed around 1450 cm⁻¹ are characteristics of Zr–O–C species. The absorption peaks 3572 cm⁻¹, correspond to the vibration of stretching and deformation due to the absorption of water. Fig. 1(d) shows the characteristic peaks for Se,Mn-HAP/ZrO₂ coating. The broad stretching peak at 3439 cm⁻¹ and a bending band at 1651 cm⁻¹ are attributed to the stretching and bending modes of water molecule. The characteristic peaks located at 1015 cm⁻¹ (ν₃) and 599 cm⁻¹ & 525 cm⁻¹ (ν₄) as well as the peaks observed at 1065 cm⁻¹ (ν₃) and 982 cm⁻¹ (ν₁) were assigned to the phosphate groups in Se,Mn-HAP.⁴⁵ Moreover, the absorption peaks at 3670 cm⁻¹ and 650 cm⁻¹ are due to the stretching and bending vibrations of OH groups of Se,Mn-HAP, respectively. The peak at 420 cm⁻¹ is attributed to the Zr–O stretching. All these peaks support for the formation of Se,Mn-HAP/ZrO₂ dual layer coating on AZ91 Mg alloy.

3.1.2 XRD analysis. The XRD patterns of HAP, Se,Mn-HAP, ZrO₂, Se,Mn-HAP/ZrO₂ dual layer coated AZ91 Mg alloy are shown in Fig. 2. The diffraction peak values (2θ) of 25.88°, 31.77°, 32.2°, 32.9°, 46.71°, 49.4° and 53.14° were assigned to HAP (Fig. 2(a)). All these diffraction peaks confirm the presence of HAP and agree well with the International Centre for Diffraction Data (ICDD card no. 09-0432). Fig. 2(b) depicts the XRD pattern obtained for the Se,Mn-HAP coated AZ91 Mg alloy which is in good agreement with the standard data for HAP. However, the peak positions deviate slightly from that of the HAP patterns and the corresponding 2θ values for Se,Mn-HAP coated AZ91 Mg alloy were observed at 25.21°, 31.12°, 31.7°, 32.1°, 45.4°, 49.1° and 52.9°, respectively and no other secondary peaks were found. The diffraction peaks of ZrO₂ coated AZ91 Mg alloy (Fig. 2(c)) are located at 2θ values of 34.4°, 31.8°, 36.2°, 47.4° and 56.7° which are in agreement with ICDD card no. 89-0511.49. Fig. 2(d) shows the XRD pattern obtained for the dual layer (Se,Mn-HAP/ZrO₂) coating on AZ91 Mg alloy. The major intense peaks observed at 2θ values of 25.21°, 31.11°, 31.6°, 32.1°, 45.5°, 49.1° and 52.8° correspond to Se,Mn-HAP. The remaining peaks observed at 2θ values of 36.2°, 47.4° and 56.7° correspond to ZrO₂. As can be seen from Fig. 2(b) and (c), the characteristic peaks of both Se,Mn-HAP and ZrO₂ are found in the XRD pattern of the Se,Mn-HAP/ZrO₂ dual layer coating (Fig. 2(d)). This evidences the formation of Se,Mn-HAP/ZrO₂ dual layer coating on AZ91 Mg alloy.

Fig. 2 XRD patterns of (a) HAP (b) Se,Mn-HAP (c) ZrO₂ and (d) Se,Mn-HAP/ZrO₂ dual layer coated AZ91 Mg alloy.

3.1.3 SEM and EDX analysis. As the morphology and nature of the coatings are known to govern the bio-physiological and mechanical properties of the coated implant materials, the coated AZ91 Mg alloy implants were subjected to detailed characterization by SEM. The surface morphologies of ZrO₂ coatings on AZ91 Mg alloy at three different time duration of 600 s, 1200 s and 1800 s for the 1 mA cm⁻² are shown in Fig. 3(a–c), respectively. Many cracks were found on the ZrO₂ coating developed at 600 s (Fig. 3(a)) on AZ91 Mg alloy. After the deposition time of 1200 s (Fig. 3(b)), uniform and mud crack free morphology of ZrO₂ could be found on the surface of the AZ91 Mg alloy. On further increasing the time to 1800 s (Fig. 3(c)), slightly agglomerated, non homogenous structure of ZrO₂ coating was observed. The elemental composition of the ZrO₂ is shown in Fig. 3(d). On comparing the morphology of the ZrO₂ coatings at different time, the coating obtained at 1200 s (Fig. 3(b)) consisted of uniform and mud-cracks free coating which is considered as an optimum for the development of a second layer of bioceramic material (HAP, Mn-HAP and Se,Mn-HAP) over the ZrO₂ coated AZ91 Mg alloy. Moreover, the morphology of ZrO₂ coating on AZ91 Mg alloy revealed a rough surface with a uniformly distributed microstructure which is believed to be having a favorable surface for the subsequent layer of bioceramic material that is to be developed over it.

Fig. 3 HRSEM images of the ZrO₂ coatings on AZ91 Mg alloy at three different deposition time (a) 5 min, (b) 10 min and (c) 15 min and (d) EDAX spectra of ZrO₂ at 10 min.

Fig. 4 shows the morphology of HAP, Mn-HAP and Se,Mn-HAP coating on ZrO₂ coated AZ91 Mg alloy. The HAP coating on ZrO₂ coated AZ91 Mg alloy (Fig. 4(a)) exhibits compact, arrangement of microstructure with few pores in between them. The surface morphology of the Mn-HAP coating on ZrO₂ coated AZ91 Mg alloy (Fig. 4(b)) exhibited the formation of compact, small leaf-like microstructure on the entire surface. Whereas, when Se is substituted into Mn-HAP coating (Fig. 4(c)), small rose flowers like structures begin to appear, thereby covering the entire surface. Fig. 4(d) gives the enlarged view of the rose flower morphology of the Mn,Se-HAP coating developed over the ZrO₂ coated AZ91 Mg alloy. It is worth noting that most of the rose flower-like structures are interconnected which should be encouraging for bone ingrowths on the surface of the Se,Mn-HAP/ZrO₂ dual layer coated AZ91 Mg alloy. The rose flower-like structures pave the way for the proliferation and attachment of cells responsible for the new bone formation. However, the porous, rose flower like structure of the Se,Mn-HAP coating on ZrO₂ coated AZ91 Mg alloy may be favorable for the bone formation by providing the essential nutrients for the growth of osseous tissue. Further, the formation of Se,Mn-HAP/ZrO₂ dual layer coatings is confirmed by the EDS mapping pattern (Fig. 5) which shows the images for the elements like Ca, O, Se, Mn, Zr and P, respectively. Thus, the homogenously distributed mineral ions can be clearly observed upon EDS mapping analysis.

Fig. 4 HRSEM images of (a) HAP (b) Se-HAP (c) Mn-HAP and (d) Se,Mn-HAP coating on ZrO₂ coated AZ91 Mg alloy.

Fig. 5 EDAX mapping analysis of Se,Mn-HAP/ZrO₂ dual layer coated AZ91 Mg alloy.

3.2 Electrochemical characterizations

3.2.1 Potentiodynamic polarization measurements. Fig. 6(a) shows the potentiodynamic polarisation curves of uncoated AZ91 Mg alloy, ZrO₂ coated, Se,Mn-HAP coated and Se,Mn-HAP/ZrO₂ dual layer coated AZ91 Mg alloy samples, respectively in SBF solution. The corrosion potential (E_corr) and corrosion current density (I_corr) that are obtained from the polarization curves are presented in Table 2. The E_corr and I_corr values of uncoated AZ91 Mg alloy was found to be as −1510 mV vs. SCE and 9.1 mA cm⁻², respectively. Whereas when coated with ZrO₂, the E_corr and I_corr values become −1205 mV vs. SCE and 1.2 mA cm⁻², respectively. While the polarization curves of the Se,Mn-HAP coated AZ91 Mg alloy sample shows the E_corr and I_corr values as −1310 mV vs. SCE and 5. mA cm⁻², the Se,Mn-HAP/ZrO₂ dual layer coated AZ91 Mg alloy sample exhibited E_corr and I_corr values as −1100 mV vs. SCE and 0.8 mA cm⁻², respectively (Fig. 6(a)). From the obtained E_corr and I_corr values it can be clear that ZrO₂ coated AZ91 Mg alloy sample were found to be nobler than that of the Se,Mn-HAP coated and uncoated AZ91 Mg alloy specimens which is owing to the uniform and compact surface coverage of the ZrO₂ on the AZ91 Mg alloy sample. The lower E_corr values of the Se,Mn-HAP coated AZ91 Mg alloy is due to the loosely arranged porous structure. So, based on these, the ZrO₂ coating can be developed as a primary layer prior to Se,Mn-HAP coating on AZ91 Mg alloy for better protection against corrosion in the SBF solution. The maximum shift of E_corr value towards the less negative direction (noble direction) is an indication that the Se,Mn-HAP/ZrO₂ dual layer coating possessed higher corrosion resistance in SBF solution, when compared to that of the uncoated, Se,Mn-HAP and the ZrO₂ coated AZ91 Mg alloy samples.

Fig. 6 (a) Potentiodynamic polarisation and (b) Nyquist plots for uncoated, Se,Mn-HAP, ZrO₂ and Se,Mn-HAP/ZrO₂ dual layer coated AZ91 Mg alloy specimens in SBF solution.

Table 2 Electrochemical parameters of the uncoated, Se,Mn-HAP, ZrO₂ and Se,Mn-HAP/ZrO₂, coated AZ91 Mg alloy in SBF solution

Sample condition	Polarisation parameters		Impedance parameters
	Ecorr (mV vs. SCE)	Icorr × 10^−5 (mA cm^−2)	Rp (Ω cm^2)
Uncoated	−1510	9.1	1020
Se,Mn-HAP	−1310	5.0	3150
ZrO2	−1205	1.2	3460
Se,Mn-HAP/ZrO2	−1100	0.8	3850

---

3.2.2 Electrochemical impedance spectroscopic studies. EIS is the most significant technique which can offer the useful information on both resistive and capacitive behavior of all the as-coated AZ91 Mg alloy specimens in the SBF solution. Fig. 6(b) shows the Nyquist plots of uncoated, ZrO₂ coated, Se,Mn-HAP coated and Se,Mn-HAP/ZrO₂ dual layer coated AZ91 Mg alloy samples. The polarisation resistance (R_p) values for the uncoated AZ91 Mg alloy sample is found as 1020 Ω cm² whereas for the Se,Mn-HAP coated AZ91 Mg alloy sample the R_p becomes 3150 Ω cm². The ZrO₂ coated AZ91 Mg alloy sample exhibited a greater R_p value (3460 Ω cm²) than that of the Se,Mn-HAP coated AZ91 Mg alloy and uncoated AZ91 Mg alloy specimens, respectively (Table 1). The R_p value for Se,Mn-HAP/ZrO₂ dual layer coated AZ91 Mg alloy sample is found to be 3850 Ω cm² which is greater than that of the ZrO₂ coated AZ91 Mg alloy specimen. For the dual layer (Se,Mn-HAP/ZrO₂) coated AZ91 Mg alloy sample, the Nyquist plot exhibits two capacitive semicircles and in which, the higher frequencies semicircle (first semicircle) can be attributed to the Se,Mn-HAP (rose flower like layer) and the second semicircle at low frequencies corresponds to the ZrO₂ layer (uniform and compact). The greater R_p values of the Se,Mn-HAP/ZrO₂ dual layer coatings are due to the superior and effective barrier of compact ZrO₂ layer formed on the AZ91 Mg alloy substrate prior to the Se,Mn-HAP coating. From this study it is well evident that, ZrO₂ serves as a protective barrier layer between the top layer (Se,Mn-HAP) and AZ91 Mg alloy and thus the Se,Mn-HAP/ZrO₂ dual layer coated AZ91 Mg alloy with higher corrosion resistance can serve as a potential material for in vivo implantation (load bearing applications).

3.3 Mechanical characterization of the Se,Mn-HAP/ZrO₂ dual layer coatings

Adhesion strength of the coating on the metallic implants is one of the important properties for in vivo implantation. Here, the adhesion strength of the Se,Mn-HAP, ZrO₂ and dual layer Se,Mn-HAP/ZrO₂ coatings on AZ91 Mg alloy (see Fig. S1(a) in ESI†) was evaluated. The adhesion strength of the Se,Mn-HAP coating on the ZrO₂ coated AZ91 Mg alloy (15.2 MPa) was higher than that of the individual coating of Se,Mn-HAP (11.6 MPa) and ZrO₂ (17.5 MPa) coated AZ91 Mg alloy, respectively. This improved adhesion strength of the as-formed dual layer coating will make it suitable for load bearing applications. The hardness of the as developed coatings follows the same pattern as observed for the adhesion strength. The Vickers micro-hardness (H_v) values for the AZ91 Mg alloy, Se,Mn-HAP, and dual layer (Se,Mn-HAP/ZrO₂) coatings on AZ91 Mg alloy, respectively are shown in (see Fig. S1(b) in ESI†). The Vicker's micro-hardness value of uncoated and Se,Mn-HAP coated AZ91 Mg alloy specimens was found to be 89.1 ± 5.2 Hv and 341.5 ± 10.2 Hv, respectively. Whereas, the Vicker's micro-hardness values (420.3 ± 12.1 Hv) obtained for the dual layer coated AZ91 Mg alloy was higher than that of the uncoated and Se,Mn-HAP coated AZ91 Mg alloy specimens. This may be attributed to the compact and uniform coating of Se,Mn-HAP over ZrO₂ coated AZ91 Mg alloy.

3.4 Biological characterization of the Se,Mn-HAP/ZrO₂ dual layer coatings

3.4.1 Antibacterial activity. S. aureus (Gram-positive) and E. coli (Gram-negative) are the general bacteria's that are found in the contaminated wound. The as developed HAP, Se-HAP, Se,Mn-HAP and Se,Mn-HAP/ZrO₂ coatings were tested against S. aureus and E. coli stains at different concentrations like 25, 50, 75, 100 and 125 (μg mL⁻¹) and compared with control (Fig. 7(a–c)). From the figure it can be well evident that, Se,Mn-HAP/ZrO₂ dual layer coating gave higher antibacterial activity against both the bacterial strains. This is because of the substitution of Se in Mn-HAP that can react with the nuclear content of bacteria and destroy them. Also some researchers argue that Zr ions are also responsible for their antibacterial activity and hence the Se,Mn-HAP/ZrO₂ dual layer coatings showed excellent anti-bacterial activity. Comparatively, the antibacterial activity of coatings against E. coli is slightly higher than that of S. aureus.^54,55

Fig. 7 Antibacterial activity of HAP, Se-HAP, Se,Mn-HAP and Se,Mn-HAP/ZrO₂ dual layer coating on AZ91 Mg alloy samples with different concentrations of 25, 50, 75, 100, and 125 against S. aureus (a), E. coli (b) and (c) photographs for the zone of inhibition of dual layer coating against S. aureus and E. coli.

The anti-bacterial activity results are well supported by the zone of inhibition observed from Fig. 7(c). The measured inhibition zones for the Se,Mn-HAP/ZrO₂ dual layer coating was found to be 12, 14, 15, 19, and 20 mm for S. aureus, and 10, 13, 16, 18, and 22 mm for E. coli respectively. From the plates it is observed that upon increasing the concentration of Se,Mn-HAP/ZrO₂, the measured inhibition zone for E. coli and S. aureus were increasing and in particular the coatings showed excellent activity against E. coli which is due to the differences in the cell wall structure. The cell wall of the Gram-positive bacteria is composed of a thick layer of peptidoglycan, consisting of linear polysaccharide chains cross-linked by short peptides thus forming more rigid structure leading to difficult penetration compared to the Gram-negative bacteria where the cell wall possesses thinner layer of peptidoglycan. Therefore, changes in the membrane structure of bacteria results in the increased anti-bacterial activity for the coatings against E. coli.⁵⁶ From the results it is well evident that the Se,Mn-HAP/ZrO₂ dual layer coating not only retards the bacterial adhesion, but also effectively kills the adhered bacteria suggesting effective and long lasting antibacterial activity against both E. coli and S. aureus.

3.4.2 In vitro cytotoxicity assessment of dual layer against human fibroblast cells. The cytotoxicity assay was used to determine the viability of human fibroblast cells on the dual layer (Se,Mn-HAP/ZrO₂) coating cultured for 1, 4 and 7 days. The percentage cell viability of dual layer coating was calculated with respect to the control for 1, 4 and 7 days and the results are shown as bar diagram in Fig. 8(a). From the figure it is well evident that, as the days of incubation is increased, the % cell viability also increased which is in good agreement with the fluorescence microscopic images as shown in Fig. 8(b–g). The dual layer coatings exhibited a substantial cell viability which is similar to that of the control group (Fig. 8(b)). At day 1, the cells distributed evenly and spread on the coated sample. On day 4, the cells got spread better on the surfaces when compared to that on day 1. Almost no dead cell can be observed from the samples. As the incubation was increased to 7 days, the number of cells also increased linearly. This is an indication that the dual layer (Se,Mn-HAP/ZrO₂) coating has nearly no cytotoxicity and indeed support cell proliferation. The dual layer coatings at 7 days of culture (Fig. 8(g)) showed the presence of more viable cells which evidence that the biocompatibility of the dual layer is not affected by the presence of mineral ions Se and Mn in the coating. It is also evident that ZrO₂ does not affect the bioactivity of the dual layer coating. Thus, the MTT assay test clearly shows that the dual layer of Se,Mn-HAP/ZrO₂ coating extensively increased the viability of cells (99.6%) which is favorable for orthopedic applications.

Fig. 8 Bar diagram (a) and optical microscopic images showing the % viability of human fibroblast cells on control (b, d and f) and Se,Mn-HAP/ZrO₂ dual layer coated AZ91 Mg alloy (c, e and g) at 1, 4 and 7 days of incubation.

3.4.3 In vitro anticancer assessment of dual layer coating against bone cancer cell lines. The in vitro anticancer effect of dual layer coating was employed using MTT assay. The major findings indicated that Se,Mn-HAP/ZrO₂ layer induced a dose dependent cytotoxicity effect on Saos-2 cells (Fig. 9(a–e)). This is the first study to report on the cytotoxicity of dual layer (Se,Mn-HAP/ZrO₂) coatings against bone cancer cell lines (Saos-2) at 1, 4 and 7 days of incubation. The cytotoxic effect of dual layer (Se,Mn-HAP/ZrO₂) coated sample against Saos-2 cells was found to be greater at 7 days of incubation. The cytotoxic nature of the dual layer is mainly due to the presence of Se in the coating. This result is supported by the optical microscopic images which show that the dual layer coating reduced the proliferation of cancerous cells at 7 days of incubation (Fig. 9(e)). These results support for the anti-cancer property of the Se,Mn-HAP/ZrO₂ dual layer coating on Mg alloy.

Fig. 9 Bar diagram (a) and optical microscopic images showing the % viability of bone cancer Saos-2 cells on (b) control and Se,Mn-HAP/ZrO₂ dual layer-coated AZ91 Mg alloy for (c) 1, (d) 4 and (e) 7 days of incubation.

3.4.4 Cell adhesion. Fig. 10 shows the SEM morphology of the cells cultured on Se,Mn-HAP/ZrO₂ dual layer coating on AZ91 Mg alloy at 1, 4 and 7 days of incubation. The morphology of HOS MG63 osteoblasts cells on Se,Mn-HAP/ZrO₂ dual layer coating showed good cell attachment thus indicating the better biocompatibility of the dual layer coated samples. Fig. 10(a) shows the presence of few osteoblasts cells attached to Se,Mn-HAP/ZrO₂ dual layer coating and it is observed that the cells were loosely spread over the surface. Fig. 10(b) presents the cell morphology of HOS MG63 osteoblasts on Se,Mn-HAP/ZrO₂ coated Mg alloy after 4 days of incubation. Cells were grown well on the surface for 4 days of incubation which indicates the presence of larger surface area for the interaction with the osteoblast cells. At day 7, the cells spread better on the entire surface than those at day 1 and day 4 of incubation. It is observed from this morphological study that the cell attachment on the dual layer coating was good for one day of incubation, but the level of cell attachment increased significantly greater as the days of incubation increased. It is found that the cells on Se,Mn-HAP/ZrO₂ coatings possessed both lamellipodial and filopodial structures while extending the incubation time to 7 days. These findings clearly indicate that Se,Mn-HAP/ZrO₂ coating promotes the growth of osteoblast cells and almost cover the entire surface of the coating. The results obtained for the different coatings indicate that the promotion of HOS MG63 cell attachment is mainly due to the presence of mineral ions (Se and Mn) in the coating. These mineral ions remarkably improved the cellular response to the dual layer coating. Hence, this cell culture result shows that Se,Mn-HAP/ZrO₂ composite coating on AZ91 Mg alloy have good biocompatibility without any toxicity.

Fig. 10 HRSEM micrographs of HOS MG63 cell adhesion growth on Se,Mn-HAP/ZrO₂ dual layer coating on AZ91 Mg alloy at (a) 1, (b) 4 and (c) 7 days of incubation.

3.4.5 In vivo behavior of the dual layer coating. The reconstruction of the damaged bone, cell behavior and bioactivity around implants in experimental animals were histologically evaluated (Fig. 11). All implants were fully removed from the bone of experimental animals using forceps after 14 and 28 days. It was most difficult to detach group II implants as compared to group I. After the removal of implants, the bone around the implant of groups II was compared with the control bones (group I) for the experimental animals. From (Fig. 11(b) and (d)) group II, it can be seen that, the bone tissues exhibited newly grown bone, hematopoietic marrow and trabecular bones without any inflammation. After 28 days of implantation, the medullar cavity in group II was fully restored. The cortical defect for group II was also restored due to formation of the neocortical bone. This explains the difficulty in detaching the group II implants from the bone using forceps. A normal neo-bone near the Se,Mn-HAP/ZrO₂ coating suggests that bone regeneration was complete. This supports that the Se,Mn-HAP/ZrO₂ coating has not caused any adverse effect or inflammatory reaction. But, the group I exhibited slightly reduced bone formation among the other coated implants. This study clearly envisioned that, the Se,Mn-HAP/ZrO₂ coating on AZ91 Mg alloy accelerates osseointegration.

Fig. 11 Histological images of group (I) (a and c) and group (II) (b and d) at 14 and 28 days of implantation against Mallory hematoxylin–eosin stain. The layer is thicker for group II (b) (Se,Mn-HAP/ZrO₂ coating) and attached to newly grown bone tissues.

4. Conclusions

In this paper, we have demonstrated the successful electrodeposition of Se,Mn-HAP/ZrO₂ dual layer on AZ91 Mg alloy. Having investigated the corrosion protection performance of all the as formed coatings, we found that the dual layer coating plays an effective role in improving the bioresistivity of AZ91 Mg alloy. This study also identified that the present of selenium in dual layer coating inhibited the cancerous osteoblast functions, while still promoting the healthy osteoblast functions. The promising results of histological observations showed good osseointegration which supports for the beneficial effect of Se,Mn-HAP/ZrO₂ dual layer coating on implant fixation in rats. All these results supports that the Se,Mn-HAP/ZrO₂ dual layer coated Mg alloy can be a promising implant material with multi-functional properties like enhanced corrosion resistance, mechanical, biocompatible and anti-cancer properties for biomedical applications. Further in vitro and in vivo characterization supports that the dual layer coated Mg alloy can be a promising candidate for biomedical applications.

Acknowledgements

One of the authors D. Gopi acknowledges the major financial support from the Department of Science and Technology, New Delhi, India (Ref. No. DST/TSG/NTS/2011/73 and Ref. No. SB/EMEQ-185/2013) and Council of Scientific and Industrial Research (CSIR, Ref. No. 01(2547)/11/EMR-II, Dated: 12.12.2011). Also, D. Gopi and L. Kavitha acknowledge the University Grants Commission (Ref. No. F. 30-1/2013 (SA-II)/RA-2012-14-NEW-SC-TAM-3240 and Ref. No. F. 30-1/2013(SA-II)/RA-2012-14-NEW-GE-TAM-3228) for the Research Awards.

References

1. Y. Zheng, H. Zhou, C. R. Dunstan, R. L. Sutherland and M. J. Seibel, J. Bone Oncology, 2013, 2, 47–57.
2. K. N. Weilbaecher, T. A. Guise and L. K. McCauley, Nat. Rev. Cancer, 2011, 11, 411–425.
3. A. Lipton, J. R. Berenson, J. J. Body, B. F. Boyce, O. S. Bruland and M. A. Carducci, Clin. Cancer Res., 2006, 12, 6209–6212.
4. P. Curtin, H. Youmb and E. Salih, Biomaterials, 2012, 33, 1065–1078.
5. M. P. Staige, A. M. Pietaka, J. Huadmaia and G. Dias, Biomaterials, 2006, 27, 1728–1734.
6. N. T. Kirkland, N. Birbilis and M. P. Staiger, Acta Biomater., 2012, 8, 925–936.
7. W. Cui, E. Beniash, E. Gawalt, Z. Xud and C. Sfeir, Acta Biomater., 2013, 9, 8650–8659.
8. H. Hornberger, S. Virtanen and A. R. Boccaccini, Acta Biomater., 2012, 8, 2442–2455.
9. F. Witte, V. Kaese, H. Haferkamp, E. Switzer, A. Meyer-Lindenberg, C. J. Wirth and H. Windhagen, Biomaterials, 2005, 26, 3557–3563.
10. M. Tomozawa and S. Hiromoto, Appl. Surf. Sci., 2011, 257, 8253–8257.
11. M. J. Wang, C. F. Li and S. K. Yen, Corros. Sci., 2013, 76, 142–153.
12. X. Chen, K. Cai, J. Fang, M. Lai, Y. Hou, J. Li, Z. Luo, Y. Hu and L. Tang, Colloids Surf., B, 2013, 103, 149–157.
13. E. Boanini, M. Gazzano and A. Bigi, Acta Biomater., 2010, 6, 1882–1894.
14. Y. Tang, H. F. Chappell, M. T. Dove, R. J. Reeder and Y. J. Lee, Biomaterials, 2009, 30, 2864–2872.
15. I. Cacciotti, A. Bianco, M. Lombardi and L. Montanaro, J. Eur. Ceram. Soc., 2009, 29, 2969–2978.
16. D. Gopi, S. Nithiya, E. Shinyjoy and L. Kavitha, Spectrochim. Acta, Part A, 2012, 92, 194–200.
17. D. Gopi, S. Nithiya, E. Shinyjoy, D. Rajeswari and L. Kavitha, Ind. Eng. Chem. Res., 2014, 53, 7660–7669.
18. D. Gopi, S. Ramya, D. Rajeswari, P. Karthikeyan and L. Kavitha, Colloids Surf., A, 2014, 451, 177–186.
19. J. Brozmanová, D. Mániková, V. Vcková and M. Chovanec, Arch. Toxicol., 2010, 84, 919.
20. P. A. Tran, L. Sarin, R. H. Hurt and T. J. Webster, J. Biomed. Mater. Res., Part A, 2010, 93, 1417–1428.
21. M. P. Rayman, Lancet, 2000, 356, 233–241.
22. M. Navarro-Alarcon and M. C. Lopez-Martinez, Sci. Total Environ., 2000, 249, 47.
23. G. F. Combs and W. P. Gray, Pharmacol. Ther., 1998, 79, 179–192.
24. C. Jiang, W. Q. Jiang, H. Ganther and L. X. Lu, Mol. Carcinog., 1999, 26, 213–225.
25. P. Tran and T. J. Webster, Int. J. Nanomed., 2008, 3, 391–396.
26. P. A. Tran, L. Sarin, R. H. Hurt and T. J. Webster, Int. J. Nanomed., 2010, 5, 351–358.
27. J. Kolmas, E. Oledzka, M. Sobczak and G. N. Jawecki, Mater. Sci. Eng., C, 2014, 39, 134–142.
28. V. Perla and T. J. Webster, J. Biomed. Mater. Res., Part A, 2005, 75, 356–364.
29. B. Bracci, P. Torricelli, S. Panzavolta, E. Boanini, R. Giardino and A. Bigi, J. Inorg. Biochem., 2009, 103, 1666–1674.
30. F. Sima, G. Socol, E. Axente, I. N. Mihailescu, L. Zdrentu, S. M. Petrescu and I. Mayer, Appl. Surf. Sci., 2007, 254, 1155–1159.
31. A. S. Hazell, P. Desjardins and R. F. Butterworth, Neurochem. Int., 1999, 35, 11–17.
32. R. M. Leach and A. M. Muenster, J. Nutr., 1962, 78, 51–56.
33. C. Paluszkiewicz, A. Slosarczyk, D. Pijocha, M. Sitarz, M. Bućko, A. Zima, A. Chróścicka and M. Lewandowska-Szumieł, J. Mol. Struct., 2010, 976, 301–309.
34. B. Bracci, P. Torricelli, S. Panzavolta, E. Boanini, R. Giardino and A. Bigi, J. Inorg. Biochem., 2009, 103, 1666–1674.
35. E. Bonnelye, A. Chabadel, F. Saltel and P. Jurdic, Bone, 2008, 42, 129–138.
36. M. J. Wang, C. F. Li and S. K. Yen, Corros. Sci., 2013, 76, 142–153.
37. L. Fu, K. A. Khor and J. P. Lim, Surf. Coat. Technol., 2000, 127, 66–75.
38. J. Wang, C. Huang, Q. Wan, Y. Chen and Y. Chao, Surf. Coat. Technol., 2010, 204, 2576–2582.
39. S. R. Witek and E. D. Butler, J. Mater. Sci. Lett., 1985, 4, 1412–1414.
40. B. W. Buczynski, M. M. Kory, R. P. Steiner, T. A. Kittinger and R. D. Ramsier, Colloids Surf., B, 2003, 30, 167–175.
41. Q. L. Ge, T. C. Lei and Y. Zhou, Mater. Sci. Technol., 1991, 7, 490–494.
42. M. Sandhyarani, N. Rameshbabu, K. Venkateswarlu and L. Rama Krishna, Surf. Coat. Technol., 2014, 238, 58–67.
43. Y. T. Liu, T. M. Leeb and T. S. Lui, Colloids Surf., B, 2013, 106, 37–45.
44. R. Quan, D. Yang, J. Yan, W. Li, X. Wu and H. Wang, Mater. Sci. Eng., C, 2009, 29, 253–260.
45. R. Jayakumar, R. Ramachandran, V. V. Divya Rani, K. P. Chennazhi, H. T. Amura and S. V. Nair, J. Biol. Macromol., 2011, 48, 336–344.
46. R. Brakel, M. S. Cune, A. J. Winkelhoff, C. Putter and J. W. Verhoeven, Clin. Oral Implants Res., 2011, 22, 571–577.
47. H. L. Huang, Y. Y. Chang, Y. C. Chen, C. L. Michael and Y. C. Chen, Thin Solid Films, 2013, 549, 108–116.
48. F. Samanipoura, M. R. Bayatia, F. G. Farda, H. R. Zargard, T. Troczynskid and A. R. Mirhabibia, Colloids Surf., B, 2011, 86, 14–20.
49. D. Gopi, S. Ramya, D. Rajeswari and L. Kavitha, Colloids Surf., B, 2013, 107, 130–136.
50. C. L. Wen, S. K. Guan, L. Peng, C. X. Ren, X. Wang and Z. H. Hu, Appl. Surf. Sci., 2009, 255, 6433–6438.
51. Y. W. Song, D. Y. Shan and E. H. Han, Mater. Lett., 2008, 62, 3276–3279.
52. D. Gopi, N. Murugan, S. Ramya and L. Kavitha, J. Mater. Chem. B, 2014, 2, 5531–5540.
53. T. Kokubo, H. Kushitani, S. Sakka, T. Kitsugi and T. Yamamuto, J. Biomed. Mater. Res., 1990, 24, 721–734.
54. Y. Zhao, M. I. Jamesh, W. K. Li, G. Wu, C. Wang, Y. Zheng, K. W. K. Yeung and P. K. Chu, Acta Biomater., 2014, 10, 544–556.
55. H. Qin, Y. Zhao, Z. An, M. Cheng, Q. Wang, T. Cheng, Q. Wang, J. Wang, Y. Jiang, X. Zhan and G. Yuan, Biomaterials, 2015, 53, 211–220.
56. S. Kaviya, J. Santhanalakshmi, B. Viswanathan, J. Muthumary and K. Srinivasan, Spectrochim. Acta, Part A, 2011, 79, 594–598.

---

Footnote

† Electronic supplementary information (ESI) available. See DOI: 10.1039/c5ra17747b

---