Synthesis and characterisation of nanostructured hardystonite coating on stainless steel for biomedical application

Abstract

Here, nanostructured hardystonite bioceramic (Ca₂ ZnSi₂ O₇) was synthesised from tetraethyl orthosilicate, zinc nitrate hexahydrate, and calcium nitrate tetrahydrate via sol–gel method, dried at 60–120°C, and finally calcinated at 1300°C. X‐ray diffraction (XRD) analysis confirmed the formation of hardystonite bioceramic. Afterwards, electrophoretic method was utilised to coat the hardystonite ceramic on 316L stainless steel (SS). Methanol solution was used as suspension solvent. The best deposition procedure was carried out by electrophoretic device in the voltage of 50 V for 5 min. XRD analysis was employed for phase characterisation and scanning electron microscopy was utilised for microstructural and morphological characterisations of the coatings. Chemical composition of the coating was evaluated by energy‐dispersive X‐ray spectroscopy. The hardystonite coating improved the corrosion resistance of the substrate, so the corrosion current density in the coated samples was less than the uncoated ones (nine times). In order to assess the bioactivity of the coating, simulated body fluid was used. The main results of the coated sample bioactivity demonstrated that the nanostructured hardystonite coating could amend the in vitro SS bioactivity. Therefore, SS coated with nanostructured hardystonite may be a promising candidate to be applied as bioactive hard tissue implants.

1 Introduction

Recently, the demand for long shelf‐life hard tissue implants has dramatically promoted with increment of the primary and revision total hip and knee arthroplasties. Furthermore, bone replacements are becoming prevailing among young people with active lifestyles [1]. Therefore, the bone implants with the increased shelf‐life and excellent stability in the osseointegration process are significant key factor for convincing the patients in clinical applications. In the past, only specific metals such as gold, silver, and copper were used as biomaterials but after 1221 [2], 316L SS along with cobalt, chromium, titanium, aluminium, and vanadium were also utilised. After that, polymeric biomaterials were developed and employed as artificial blood vessels, contact lenses, and surgical sutures [2]. SS is one of the most popular materials used as implants due to relatively low cost, facile manufacturing, and acceptable corrosion resistance [3, 4]. SS are susceptible of localised corrosion in long‐term applications in the body where there are biological effects. They can release iron, chromium, nickel, and molybdenum ions that aggregate around the implant tissues or transmit to other areas [5, 6]. SS corrosion products in concentrations higher than a specific limit can damage the normal behaviour of bone marrow cells proliferation and growth [7]. Therefore, coating of metallic implants is one of the suggested solutions for reduction in metal ion release [6, 7, 8].

Bioactive ceramics are used as coatings on clamps with high strength and toughness such as implants to repair the bones, joints, and teeth in injuries. In addition, they are employed as pyrolytic carbon coatings in heart valves and radioactive glasses are used for treatment of certain tumours. They are also utilised as coatings on metallic substrates. Among all of the bioactive materials, the most proper biocompatibility behaviour is belonged to hydroxyapatite (HA) bioceramics [9] and bioglasses [10, 11] considered as a group of calcium phosphate compounds which attach to graft tissue in a short time [12, 13]. Specific compounds of bioactive glasses including P₂ O₅, SiO₂, and CaO are able to attach to the soft and hard tissues without interference of fibrous layers. Different applications were reported for bioglass and HA powders such as orthopaedic and dental implants [12, 14]. Although HA has low mechanical properties, its unique biological properties caused to be considered for development of its properties rather than to be replaced with other materials. Calcium silicates such as CaSiO₃ and Ca₂ SiO₄ among conventional ceramic materials can release calcium and silicate ions approved to influence on the bone replacement and bone repair in natural environment (in living creature's body – in vivo) and out of natural environment (in vitro) [15, 16]. Ceramic composites containing zinc are providing new opportunities in bone tissue engineering studies. Ito et al. reported that zinc‐tricalcium phosphate (Zn‐TCP) improved cell proliferation. Zinc aluminate ceramics also had a stimulation effect on osteoblast cells and improved the cell differentiation and mineral formation process [17]. In another study, Wu et al. added zinc to calcium silicate and synthesised the hardystonite ($\mathrm{C}{\mathrm{a}}_2\mathrm{ZnS}{\mathrm{i}}_2{\mathrm{O}}_7$) that owned higher mechanical properties than HA [17]. Chemical stability of hardystonite was higher than calcium silicate ceramics as the presence of zinc ions that affected on the ceramic hardness [18]. Hardystonite as a calcium silicate ceramic could induce the biological fixation and tissue growth in the implant–tissue interface [19, 20]. Moreover, hardystonite was more chemically stable than other calcium silicate‐based ceramics such as $\mathrm{C}{\mathrm{a}}_2\mathrm{Si}{\mathrm{O}}_4$ and $\mathrm{CaMgS}{\mathrm{i}}_2{\mathrm{O}}_6$ [20, 21]. Furthermore, hardystonite ceramics could improve the proliferation and differentiation of bone marrow stem cells along with proper ability to bone repair induction (osteoinduction) [19].

It should be noted that a major drawback of the calcium‐silicate‐based bioceramics was their high dissolution rate, which directed to an enhanced pH amount in the setting that could be toxic for tissue cells [22]. Moreover, the calcium‐silicate could not help to proliferation of human bone cell [23]. One of the best approaches to address these drawbacks was combining ZnO into CaSiO₃ known as an important trace species that showed a key role in the metabolism process of bone [20]. Yagamuchi et al. investigated the stimulatory action of Zn on bone protein, bone formation, and alkaline phosphatase (ALP) activity in vitro and in vivo [17]. In the other study, Ito et al. showed that Zn‐substituted tricalcium phosphate improved cell proliferation [20]. Wu et al. combined Zn into calcium silicate and prepared hardystonite (Ca₂ ZnSi₂ O₇, or 2CaO·ZnO·2SiO₂, later referred to as ‘HST’) which had better mechanical strength than HA [24]. Hardystonite bioceramics were followed as promising biomaterials for hard tissue replacements and bioactive coatings. Actually, hardystonite bioceramics have exposed to stimulate osteogenic differentiation and to enhance the proliferation rate of mesenchymal stem cells [24]. Moreover, these evidences signified that hardystonite bioceramics were osteoconducive to both types of bone cells (osteoblast‐like cells and osteoclasts) that enhanced cell proliferation and differentiation and induced expression of ALP, osteocalcin, and collagen type I [22, 24, 25].

The main purpose of this research was to improve the in vitro bioactivity of the 316L SS‐based implants by a nanostructured hardystonite coating on the SS through electrophoretic deposition. Hardystonite ceramic was synthesised and prepared by calcination at 1300°C. Hardystonite ceramic was coated on 316L SS by electrophoretic method. Apatite formation on the coated hardystonite was examined by simulated body fluid (SBF) immersion test to evaluate its bioactivity.

2 Experimental

2.1 Preparation of hardystonite powder

Hardystonite powder was prepared by the use of tetraethyl orthosilicate (TEOS), zinc nitrate hexahydrate, and calcium nitrate tetrahydrate as reagents via sol–gel method. In summary, TEOS was mixed with water and 1 M HNO₃. Then, the mixture was hydrolysed by shaking for 30 min. Zinc nitrate hexahydrate and calcium nitrate tetrahydrate were added to the solution. The agitation of reactants was continued at ambient temperature for 5 h. In the next stage, the solution was held at 60°C for 24 h and then dried at 120°C for 48 h to obtain the dried gel. The resulted dried gel was milled and sieved, and then transferred to a corundum furnace and calcinated at 1300°C for 3 h to obtain hardystonite nanopowder. X‐ray diffraction (XRD) device (Bruker, model D8 Advance) was utilised in order to confirm the ceramic crystalline structure and phase analysis. Scanning electron microscopy (SEM) (TESCAN, model Vega‐3) was employed to evaluate the microstructure, morphology, and size of synthesised particles. Energy‐dispersive X‐ray spectroscopy (EDX) was used to analyse the formed phases.

2.2 Hardystonite ceramic coating on the 316L SS

For the first time, Ducheyne et al. [26] introduced the electrophoresis deposition (EPD) for HA coating on metallic implants. Currently, this method has received a lot of attention. Advantages of the electrophoretic technique were including facile preparation, coating thickness controlling, homogeneity, low‐temperature process, low‐cost equipment, possibility of complex shape coatings, ability of thick composite films, high purity deposition, no phase transformation during coating, and biomedical applications [27, 28]. EPD is resulted from moving charged particles to cathode in liquid suspension by applying electric field. Then, the deposition is carried out by particle coalescence [29, 30]. Charged particles can be deposited on the surface of anode and cathode based on their charge [31, 32]. Coatings with thicknesses <1 μm and >500 μm are producible by EPD. In fact, thickness is controlled by changing the time and applied voltage parameters (in the same way), particle size, morphology, and distribution accompanied with changing the dielectric constant of suspension environment [31, 33]. A secondary sintering is essential in order to condense the green deposition and remove the porosities [12, 33, 34]. Since the presence of oxygen can cause oxidation of metal–ceramic interface and weak adhesion of the coating to metallic substrate, sintering requires the use of oxygen‐free furnaces [32, 35].

Surface preparation of 316L SS samples was carried out for better adhesion. Samples surfaces were first ground by 60, 100, 220, 400, and 600 sandpapers and then washed with acetone solution in ultrasonic cleaner for 10 min.

To prepare hardystonite suspension, methanol was purchased from Merck (Germany) and used as solvent. About 70 ml of the solution was poured into a beaker and then 3 g of hardystonite powder was added to the solution. Afterwards, the solution was stirred at room temperature for 24 h to gain a homogeneous suspension. In the next step, suspension was placed in the ultrasonic stirrer for 30 min. Then, electrophoretic device was utilised for coating, graphite, and 316L SS were used as anode and cathode, respectively. During the coating process, the distance between the anode and cathode was 1 cm. In order to gain a uniform coating, time and voltage were selected 5 min and 50 V, respectively. To increase the coating adhesion to the substrate, specimens were sintered at 800°C for 2 h in neutral atmosphere (argon) that was consistent with previous studies [36] (Fig. 1).

Fig. 1.

Schematic representation of the coated hardystonite on the substrate

2.3 Bioactivity evaluation of hardystonite bioceramic coating

To evaluate the apatite formation ability of hardystonite bioceramic coating, immersion test was performed in SBF solution. Whole surface area of every sample was coated with lacquer except 1 cm² that was left uncoated. Then coated samples were inserted in a beaker and 10 ml of SBF solution was poured on it. The beaker was placed in a warm water bath with constant temperature (T  = 37 ± 0.1°C). Specimens were taken out after passing 3, 7, and 14 days and dried in room temperature. To examine the quality, morphology, and structure of the applied coatings, SEM was employed. The ion concentration of SBF in comparison with those in human blood plasma was exposed in Table 1. The bioceramic coatings were immersed in SBF at 36.5°C for 3, 7, and 14 days. The coating surface area/solution volume of SBF was calculated using the equation, $V_{\mathrm{s}}=S_{\mathrm{a}}/10$, where V _s is the volume of SBF (ml) and S _a the apparent surface area of samples (mm²).

Table 1.

Ion concentration of SBF and human blood plasma

Ion	Blood plasma, mM	SBF, mM
Na+	142.0	142.0
K+	5.0	5.0
Mg2+	1.5	1.5
Ca2+	2.5	2.5
Cl−	103.0	147.8
HCO3 −	27.0	4.2
HPO4 2−	1.0	1.0
SO4 2−	0.5	0.5
PH	7.2–7.4	7.40

3 Results and discussion

3.1 Microstructure of hardystonite powder

According to observed peaks in Fig. 2, it can be concluded that the dominant formed phase was calcium zinc silicate ($\mathrm{C}{\mathrm{a}}_2\mathrm{ZnS}{\mathrm{i}}_2{\mathrm{O}}_7$). There are hardystonite phase peaks that indicated the hardystonite crystalline phase with high purity. The strongest hardystonite peak was observed at 2θ  = 31.3106°, which is the main characteristic peak for hardystonite, according to JCPDS card No. 01‐075‐0916. Four strong diffraction of planes (211), (201), (310), and (201) were also detectable for the hardystonite ceramic with tetragonal crystal system. Interpretation of hardystonite XRD peak based on Scherrer's equation showed that hardystonite crystallite size was almost 78 nm that was consistent with the earlier investigations [36, 37]. Hardystonite elemental analysis is presented in Figs. 3 and 4. Results confirmed the presence of Zn, Si, and Ca basic elements with the intended amounts in the final composition of hardystonite before and after application of hardystonite coating on 316L SS.

Fig. 2.

XRD patterns of hardystonite bioceramic powders prepared by sol–gel method and sintered at 1300°C for 3 h

Fig. 3.

Microstructure of the synthesised bioceramic

(a) SEM picture, (b) Elemental analysis (EDX) of the hardystonite powder

Fig. 4.

Microstructure of the coated bioceramic

(a) SEM picture, (b) Elemental analysis (EDX), (c) Cross‐section illustration of hardystonite coated on 316L SS at the sample 50 V and 5 min

SEM micrographs of Fig. 3 demonstrated the morphology and particle size of hardystonite ceramics. It was clear that nanoparticles were either in very fine sizes (<100 nm) or in coarse agglomerates that was consistent with the previous studies [36, 38, 39]. EDX spectra of the powder are shown in Fig. 3 b. Moreover, the EDX spectra of the powder are detailed in the inset of Fig. 3 b. Based on the results of the EDX and XRD of the powder, the bioceramic was successfully synthesised.

3.2 Studying the structure of coated hardystonite

Obtaining an integrated, uniform, and defectless coating is one the most important purpose of electrophoretic method. SEM micrograph of the hardystonite coating with 5000× magnitude is presented in Fig. 4. It was obvious that hardystonite coating established by electrophoretic method was a crack‐less continuous coating. It was mentioned in other studies that such coating should provide high corrosion resistance and proper biocompatibility [40]. EDX spectra of the coated bioceramic are exposed in Fig. 4 b. Also, data extracted from the EDS spectra is summarised in the inset of Fig. 4 b. According to the main results of the EDX and XRD, the coated bioceramic on the 316L SS was successfully established. Fig. 4 c shows a cross‐section of the hardystonite coating on the 316L SS. The optimal thickness was almost 14 µm that exhibited a suitable crackles and bioactive coating on the SS.

3.3 Bioactivity assay by SBF solution immersion test

The soaking test was established to scrutinise the in vitro bioactivity of the coatings. The SEM and EDX images of the coatings after soaking in the SBF solution for 3, 7, and 14 days are shown in Figs. 5, 6, 7. As shown in the figures, the surfaces of the coatings were cracked (clusters of white particles), which were increasing by the time of immersion. The soaked sample in SBF after 14 days was contained more cracks and pits and less white particle precipitations than others.

Fig. 5.

Microstructure of the Samples

(a) SEM photograph, (b) EDX spectra of sample 1 immersed in the solution, SBF, for 3 days

Fig. 6.

Microstructure of the Samples

(a) SEM photograph, (b) EDX spectra of sample 2 immersed in the solution, SBF, for 7 days

Fig. 7.

Microstructure of the Samples

(a) SEM photograph, (b) EDX spectra of sample 3 immersed in the solution, SBF, for 14 days

By getting EDX analyses, the cauliflower‐like structure in the surfaces of the samples exposed the existence of calcium phosphate [41].

Figs. 5 b, 6 b, and 7 b reveal the elemental analyses of the coated samples after immersion in SBF solution. Regarding the figures, Fig. 5 shows no significant change in coating morphology (immersing in SBF solution for 3 days).

On the other hand, cauliflower‐shaped apatite was nucleated on the sample immersed in SBF solution for 7 days (Fig. 6). It could be seen from Fig. 7, cauliflower‐shaped apatite was grown completely and also was closer to the desirable structure. Based on the figures, white and light grey micro/nano species were mainly exhibited Ca, O, Mg, Si, and P, demonstrating a formed layer of apatite on the surfaces.

Results displayed that hardystonite coating formed apatite in the SBF solution; however, the amount of formed apatite was insignificant in initial time intervals. Although the employed method of SBF solution was an effective way for examination of bioactivity and determination of formed HA on the ceramic surfaces, the validation of this method was limited to a group of bioceramics. Silicate‐based bioceramic such as akermanite, maronite, and diopside showed excellent bioactivity, while phosphate, carbonate, and sulphate‐based bioceramics did not display any apatite formation on their surfaces in SBF [42, 43]. The comparison of cauliflower‐shaped apatite and the calcium phosphate extracted from the natural bone demonstrated the superior osteoconductivity [41] and bioactivity (this research).

According to the XRD pattern of the coated sample before and after immersing in SBF, there are two diffraction peaks of HA and hardystonite (Fig. 8). The samples in Figs. 8 a and b showed pure hardystonite that were related to coated sample before and after 3 days soaking in SBF, demonstrating formed HA on the sample was negligible in 3 days. After 7 and 14 days of soaking (Figs. 8 c and d), there are some peaks related to HA, indicating the bioactivity was time‐dependent for the coated samples.

Fig. 8.

XRD patterns of coated sample before and after soaking in SBF. HT means hardystonite

(a) 0, (b) 3, (c) 7 and (d) 14 days of immersing in the SBF

The tafel polarisation diagram of 316L SS specimens with and without hardystonite coating in a Ringer solution at 37 ± 1°C are shown in Fig. 9. The average values of the corrosion current and the corrosion potential of non‐coated SS in a Ringer solution are presented in Table 2 (determined by the polarisation curves and the tafel extrapolation method). According to Table 2 and Fig. 10, 316L SS without coating showed a higher corrosion density in the Ringer solution (I _corr  = 0.34). The corrosion density of the coated sample decreased in the ringer solution.

Fig. 9.

Toughened tuff polarisation test example of St‐1 = 316 L SS without coating, St‐2 = 316 L SS coating with 30 V in 3 min, St‐3 = 316 L SS coating with 30 V in 3 min, St‐4 = 316 L SS coating with 50 V in 5 min, and St‐5 = 316 L SS coating with 50 V in 3 min

Table 2.

Average potential and corrosion density in a Ringer solution at 37°C

CR, mpy	i corr, μA/cm2	R p, kΩ cm2	E corr SCE, V	β c	β a	Samples
0.0038	0.34	82.52	−0.14	0.14	0.12	St‐1
0.0030	0.27	120.61	−0.55	0.15	0.15	St‐2
0.0024	0.22	185.42	−0.36	0.21	0.17	St‐3
0.0004	0.04	786.08	−0.37	0.15	0.14	St‐4
0.0006	0.05	534.95	−0.30	0.14	0.11	St‐5

Fig. 10.

Comparison of the corrosion rate for St‐1 = 316 L SS without coating, St‐2 = 316 L SS coating with 30 V in 3 min, St‐3 = 316 L SS coating with 30 V in 3 min, St‐4 = 316 L SS coating with 50 V in 5 min, and St‐5 = 316 L SS coating with 50 V in 3 min

The comparison of the curves showed that by applying the hardystonite coating, the potential leaded to more positive values and the corrosion current density decreased. The potential of the coated sample (SCE −0.37 V) increased compared to the uncoated sample (SCE −0.14 V).

According to the reported results in Table 2, the corrosion rate of St‐1 has been calculated higher than other samples. The corrosion rate of St‐4 also was lower than control. Fig. 10 shows the comparison of the corrosion rate of coated and uncoated 316 L SS.

The value of R _p (polarisation resistance) represented the resistance of the system to corrosion. The results showed that R _p of St‐4 and St‐5 were 9.5 and 6 times more than the St‐1 sample, respectively. Fig. 11 shows a better comparison of R _p of the samples.

Fig. 11.

Polarisation resistance of the coatings, St‐1 = 316 L SS without coating, St‐2 = 316 L SS coating with 30 V in 3 min, St‐3 = 316 L SS coating with 30 V in 3 min, St‐4 = 316 L SS coating with 50 V in 5 min, and St‐5 = 316 L SS coating with 50 V in 3 min

It seems that the formation of porous coatings having high amounts of pores caused pitting and local corrosion of the underlying substrate. With the increment of voltage, the amount of deposited ceramic was increased on the substrate and consequently, the thickness of the coating was increased.

3.3.1 Electrochemical impedance spectroscopy

Fig. 12 shows a typical extension for the evaluation of electrochemical impedance data titled Nyquist chart. The Nyquist plot of the coatings in the Ringer solution at 37 ± 0.1°C is shown in Fig. 12.

Fig. 12.

Nyquist plots of the samples

Fig. 13 shows the Bode diagrams for the same data of the Nyquist plot in Fig. 12.

Fig. 13.

Bode diagram of different samples

In the above diagrams, all specimens except for the bare sample were two time constants. The analysis of single time constant (STC) diagrams was used to evaluate the electrochemical equivalent circuit accurately.

The time could be assumed to be equivalent to the electrochemical circuit (Fig. 14).

Fig. 14.

Equivalent circuit of STC

In this circuit, R _s is the solution resistance, R _ct is the charge transfer resistance, and Q is called a non‐ideal capacitor or constant phase element (CPE). The CPE was used instead of ideal capacitor because the created double‐layer planes between the electrode surface and electrolyte were not exactly parallel and the surface of the electrode was also rough and heterogeneous. The difference between these two elements was evident in the impedance formula. The impedance of capacitor is defined as Z  = a /jωC and this value for CPE is Z  = 1/(Y ₀ jω)n. In this formula, C is the capacitance, ω is the phase, Y ₀ is the admittance (reverse of impedance and equivalent of the capacitance parameter in ideal capacitor), and j is the imaginary term $\sqrt{-1}$. As could be seen, the difference between these two phrases is only the n exponent which is the value between zero and one. Its zero value refers to the ideal resistance and its one value relates to the ideal capacitor. By calculating the total impedance of this circuit, the following formula is obtained:

$${\left(\frac{R_{\mathrm{ct}}}{2}\right)}^2=Z^{\mathrm{\prime }\mathrm{\prime }2}+{\left(Z^{\mathrm{\prime }}-\left(\frac{R_{\mathrm{ct}}}{2}+R_{\mathrm{s}}\right)\right)}^2$$

This equation is a circle with radius of $R_{\mathrm{ct}}/2$ and the distance of its centre from the origin is equal to $(R_{\mathrm{ct}}/2)+R_{\mathrm{s}}$. Therefore, it can be found that with increasing the diameter of the semicircle in the Nyquist plot, the value of charge transfer resistance (R _ct) increases against corrosion. Typically, the presence of a coating on the metal surface makes the system two times constant. Different equivalent circuits can be considered for two time constant systems, which according to the accuracy of the fit, the following equivalent circuit is used for the samples in this study. This circuit consists of three resistors (respectively, from the left solution resistance, coating resistance, and charge transfer resistance) and two CPEs (for the coating and double layer) (Fig. 15).

Fig. 15.

Equivalent circuit of two time constant diagrams

In this case, as before, at the highest frequency, the impedance is equal to the solution resistance and at the lowest frequency is equal to the total resistance of the system (sum of solution resistance, coating resistance, and charge transfer resistance). Therefore for the Nyquist plot, as the diameter of the semicircles increases, the corrosion resistance of the system increases. Thus, from the Nyquist plots, it can be concluded that the 50 V–5 min sample had the largest diameter and consequently, the highest corrosion resistance.

Equalisation of the measured samples with electrochemical equivalent circuit was performed using Zsimps software.

The values of the parameters obtained from this equation are given in Table 3.

Table 3.

Results of the fitting impedance test results on the electrochemical equivalent circuit

Sample	R s, Ω cm2	R c, Ω cm2	Q c, S.s n /cm2	n	Q dl, F/cm2	n	R dl, Ω cm2
bare	18.04	—	—	—	4.8 × 10−4	0.66	3467
30 V–3 min	11.63	4354	4.73 × 10−5	0.59	1.33 × 10−5	0.69	340.7
30 V–5 min	16.69	4661	3.88 × 10−5	0.59	3.93 × 10−6	1	396.75
50 V–3 min	18.32	14,665	5.44 × 10−5	0.59	2.41 × 10−6	0.91	406.3
50 V–5 min	22.31	19,354	2.67 × 10−5	0.59	1.91 × 10−6	0.95	1175.06

As shown in Table 3, the highest coating resistance and charge transfer resistance referred to the 50 V–5 min sample.

The results were in good agreement with the results of the polarisation test. If the corrosion resistance is assumed the sum of the two coating and charge transfer resistances, the results of the sum of these resistances are shown in Fig. 16.

Fig. 16.

Result of the samples from the EIS test

As shown in Fig. 16, increasing the voltage from 30 to 50 V leaded to the increased corrosion resistance of the samples. From the Bode diagrams, the two impedance parameters at the lowest frequency and the phase angle at the highest frequency could express corrosion resistance of the coating and its preventive properties. The 50 V–5 min sample and then the 50 V–3 min sample had the highest impedance at the lowest frequency and highest phase angle at the highest frequency. Moreover, the uncoated sample had the least corrosion resistance among the samples. The results of the impedance test confirmed the results of the polarisation test.

4 Conclusion

In this research, the nanostructured hardystonite coating was established on the 316L SS using the deposition of electrophoretic. The in vitro bioactivity behaviours of hardystonite‐coated samples were studied by the SBF immersion test. Briefly, the subsequent conclusion could be obtained:

• Deposition and uniform coatings without cracks were obtained by this method.

• The optimum conditions for hardystonite heat treatment to achieve the best substrate were calcination in 800°C for 2 h.

• Synthesised hardystonite biocompatibility was in acceptable level in a way that after 14 days of immersion in SBF solution, HA layer was completely formed on it.