Topography,wetting,and corrosion responses of electrodeposited hydroxyapatite and fluoridated hydroxyapatite on magnesium

Abstract

In this study, different types of calcium-phosphate phases were coated on NaOH pre-treated pure magnesium. The coating was applied by electrodeposition method in order to provide higher corrosion resistance and improve biocompatibility for magnesium. Thickness, surface morphology and topography of the coatings were analyzed using optical, scanning electron and atomic-force microscopies, respectively. Composition and chemical bonding, crystalline structures and wettability of the coatings were characterized using energy-dispersive and attenuated total reflectance–Fourier transform infrared spectroscopies, grazing incidence X-ray diffraction and contact angle measurement, respectively. Degradation behavior of the coated specimens was also investigated by potentiodynamic polarization and immersion tests. The experiments proved the presence of a porous coating dominated by dicalcium-phosphate dehydrate on the specimens. It was also verified that the developed hydroxyapatite was crystallized by alkali post-treatment. Addition of supplemental fluoride to the coating electrolyte resulted in stable and highly crystallized structures of fluoridated hydroxyapatite. The coatings were found effective to improve biocompatibility combined with corrosion resistance of the specimens. Noticeably, the fluoride supplemented layer was efficient in lowering corrosion rate and increasing surface roughness of the specimens compared to hydroxyapatite and dicalcium-phosphate dehydrates layers.

Keywords

Magnesium biodegradable implant fluoridated hydroxyapatite electrodeposition corrosion

1. Introduction

The first trials of utilizing magnesium and its alloys as a biomaterial in orthopaedic and vascular applications took place in the late 1800s [1]. Thanks to the similar mechanical properties of magnesium and human’s natural bone; magnesium has become a desirable biodegradable material for load-bearing implants applications [2]. In addition, magnesium has non-toxic degradation products (oxides or hydroxides) and is biocompatible [3]. However, one of the critical limitations for using magnesium is the high degradation rate in physiological environments; leading to the elevation of body fluid pH as a result of hydrogen evolution [4]. The pH elevation reduces magnesium’s biocompatibility by deterring the cells adhesion to the surface of implant [5].

Various methods have been employed to modulate magnesium degradation rate in physiological environments. Alloying has been studied as a solution for decreasing magnesium’s degradation rate [6], however, the results are not promising. This is due to the formation of cathodic sites in microstructure of magnesium causing drastic electrochemical corrosion [7], and forcing alloying extent in very small range. Moreover, the potential toxicity of the alloying elements jeopardises the magnesium alloys’ biocompatibility [8]. Although, alloying improves mechanical strength of magnesium, pure magnesium’s mechanical strength meets a wide range of orthopaedic applications’ requirements [9,10]. Thus, pure magnesium could be an ideal option for moderately degrading orthopaedic implants. On the other hand, surface treatment has been reported to be a viable solution for the degradation concerns of magnesium in physiological environments [11].

A bone mimicking material would be a promising choice for a biocompatible and protective coating to control the degradation rate of magnesium implants [12]. Because bone mainly consists of calcium-phosphate (Ca–P) minerals; proposed coating materials for orthopaedic implants should be composed of the same minerals [13,14]. Several techniques, including sol–gel [15], plasma spraying [16], electrochemical deposition [17], electrophoretic deposition [18,19], and electrodeposition [8] have been used to coat Ca–P on biocompatible metals. Among these methods, the electrodeposition is a relatively simple and economical technique which is extensively used for treating on magnesium alloys [20]. Electrodeposition process have a diversity of benefits such as control of the thickness, homogeneity of the coating, capability of coating complex-shaped components, and low processing temperature [21]. It has been reported that sodium hydroxide (NaOH) pre-treatment helps electrodeposition of Ca–P compositions to be easily transformed to bone like apatite [22]. In addition, this pre-treatment causes magnesium’s surface passivation and consequently to increases cell adhesion and biocompatibility [5].

It has been reported that the electrodeposited Ca–P compositions could contain several phases including dicalcium-phosphate dehydrate (DCPD, CaHPO₄ · 2H₂O) and hydroxyapatite (HA, Ca₁₀(PO₄)₆(OH)₂) phases. Due to unrestricted solubility of DCPD phase in physiological environments, it should be transformed to HA phase, which is known to be more stable [23]. One method to enhance the Ca–P compositions’ stability is introducing fluoride ion (F⁻) to the structure; F⁻ would be embodied to coalesce fluoridated hydroxyapatite structure (FHA, (Ca₅(PO₄)₃(OH)_1−xF_x)) [24]. FHA has higher stability and lower solubility in physiological environments in comparison with HA [25]. Trace of F⁻ can help mineralization and crystallization of Ca–P compositions to incorporate in bone regeneration [26].

This study aims to comprehensively investigate the morphology and corrosion, topography and wettability responses of HA and FHA coatings on a NaOH pre-treated pure magnesium. Lacking the mechanical strength required for heavy load bearing applications, pure magnesium is undeniably suitable – due to incomparable low price and fabrication simplicity – to be used for delicate applications. Typical delicate applications could include fixation devices for metacarpals and phalanges injuries.

2. Experimental methods and materials

2.1. Specimen and coating electrolyte preparation

Magnesium specimens with dimension of 20 mm × 10 mm × 5 mm were cut from a commercially pure magnesium ingot (99.9% purity, with iron, copper, and chromium impurities of less than 0.009 wt.%) and ground using silicon carbide sand papers successively up to 1200 grit. Afterwards, the specimens were ultrasonically cleaned in ethanol for 15 min and rinsed with deionised water then dried by blowing hot air. Next, the specimens were soaked in 1M NaOH solution at ambient temperature for 30 min (NaOH pre-treatment) and rinsed by distilled water. Two different coating electrolytes were prepared through this study to coat the specimens. One, named DCPD-electrolyte, was prepared by blending 0.25 mol/l of ammonium dihydrogen phosphate (NH₄H₂PO₄) and 0.42 mol/l of calcium nitrate tetrahydrate (Ca(NO₃)₂ · 4H₂O) in deionised water to develop HA coating. The other one to deposit FHA coating, FHA-electrolyte, was prepared by mixing 0.25 mol/l of NH₄H₂PO₄, 0.42 mol/l of Ca(NO₃) · 4H₂O, 1 mol/l of sodium nitrate (NaNO₃) and 0.02 mol/l of sodium fluoride (NaF) in deionised water. In both electrolytes, Ca/P molar ratio was kept steady at 1.67 to resemble HA, and pH was adjusted to 4 by the addition of minimal amount of either nitric acid (HNO₃) or ammonium hydroxide (NH₄OH) [8].

2.2. Coating deposition

The coatings were deposited via an electrodeposition method in which a graphite electrode and the specimens were used as anode and cathodes, respectively. Deposition was performed using a current density of 1 mA/cm² for 60 min at ambient temperature while the electrolytes were rendered homogeneous by continuous stirring at 100 rpm. After DCPD deposition, the coated specimens were rinsed by distilled water and dried by hot air. The DCPD-coated specimens were then post-treated by immersing in 0.25 mol/l NaOH solution in a water bath at 80°C for 2 hours to transform the coating to HA structure. The post-treated specimens were rinsed by distilled water and then dried at 40°C for 1 hour. On the other hand, after deposition of FHA, the coated specimens were merely rinsed by distilled water and dried at 40°C for one hour.

2.3. Coating characterization

Surface morphology of the coated specimens was studied using a scanning electron microscope (SEM, JEOL 6390) equipped with energy-dispersive spectroscopy (EDS, PHILIPS XL40). Three different areas were analyzed and average values were reported to provide reliable results for the EDS analysis. The coated specimens were inspected visually and their cross-sections were further checked under optical microscope (Nikon Microphot-FXL) to assure the presence of consistent coatings. In addition, the coating thickness of the specimens was measured at 5 random spots using the optical microscope. Phase composition of the coatings were detected by Grazing Incidence X-ray diffraction (GIXRD, PANalytical X’Pert Pro, using Cu K_α radiation, $λ = 1.5406 Å$ ). Crystallinity degree ( $X_{c}$ ) of the HA phase formed on the surface of the specimens was calculated using Eq. (1), in which β(002) is the full width of the peak at half maximum (FWHM (degree)) of plane (002) reflection and K is a constant chosen as 0.24 [25,27] $\begin{matrix} (1) & β (002) \sqrt[3]{X_{c}} = K . \end{matrix}$

Attenuated Total Reflectance-Fourier Transform Infrared (ATR-FTIR) spectroscopy (Thermo Scientific iD5 Diamond ATR 133, Nicolet iS5 FTIR Spectrometer) was utilized to determine the chemical bonding of the coatings in a spectral range of 4000–600 cm⁻¹. Topography of the coated and uncoated specimens’ surfaces was evaluated by atomic-force microscopy (AFM, NanoScope IV; Digital Instruments) in the tapping mode. The Root Mean Square (RMS) surface roughness, defined as the height fluctuations in the given area, was quantitatively measured from 100 µm² of the specimens’ surface. Surface roughness measurement was carried out on three random points of corresponding specimens’ surface to obtain reproducible results. Surface wettability was evaluated by a Water Droplet Contact Angle-based test (VCA optima, AST Drop Contact Angle) according to ASTM D7334-08 standard. A $1 \pm 0.1 µ l$ drop of water was employed to measure the contact angles in a specified time of 5 s. The contact angle tests were randomly repeated for five times to obtain reliable results and the mean of the results is reported.

2.4. Electrochemical corrosion tests

Corrosion behavior of the uncoated and coated magnesium specimens was determined using potentiodynamic polarization test (PARSTAT 2263 potentiostat/galvanostat) based on ASTM G-5 standard. A three-electrode cell, including reference electrode (saturated calomel electrode (SCE)), counter electrode (graphite rod) and working electrode (the specimens) were used for the electrochemical measurements. Corrosion evaluations were carried out in 500 ml of Kokubo simulated body fluid (SBF), set to pH = 7.4 at $37 \pm 0.5 ° C$ , via constant scan rate of 1 mV/s. The measurements were repeated three times in order to eliminate deviated results. Table 1 compares the chemical composition of Kokubo SBF and human blood plasma. In order to attain supplementary information on corrosion behavior of the uncoated and coated specimens, immersion test was conducted for 14 days based on ASTM G31-72 standard. This test was performed in Kokubo SBF with pH 7.4 at $37 \pm 0.5 °$ . To complement the required information, pH variation of the SBF was monitored throughout the test.

Table 1
Chemical composition of Kokubo SBF and human blood plasma [28]

Ion concentration (mmol/L)

Na⁺ K⁺ Ca²⁺ Mg²⁺ HCO³⁻ Cl⁻ HPO₄ ²⁻ SO₄ ²⁻

Kokubo SBF 142.0 5.0 2.5 1.5 4.2 147.8 1.0 0.5

Plasma solution 142.0 5.0 2.5 1.5 27.5 103.0 1.0 0.5

	Ion concentration (mmol/L)
Kokubo SBF	142.0	5.0	2.5	1.5	4.2	147.8	1.0	0.5
Plasma solution	142.0	5.0	2.5	1.5	27.5	103.0	1.0	0.5

3. Results and discussion

3.1. Crystalline and chemical characterization

Figure 1 represents GIXRD patterns of the NaOH pre-treated and DCPD, HA, and FHA-coated specimens. The Mg(OH)₂ and Mg phases were detected on the NaOH pre-treated specimen (Fig. 1(a)). There were Mg(OH)₂ peaks at 38.268°, 50.976° and 58.763° of 2θ, emerging from the NaOH pre-treatment (Eq. (2)). Mg peaks were detected at 32.173°, 34.385°, 36.604°, 47.808° and 57.363° of 2θ. Mg(OH)₂, Mg and distinct DCPD peaks along with apatite peaks were detected on DCPD-coated specimen (Fig. 1(b)). The DCPD-electrolyte was comprised of calcium ions (Ca²⁺) and acidic phosphates (2H₂PO₄ ⁻ and 2HPO₄ ⁻) which underwent reduction reactions (Eqs (3) and (4)). Simultaneously, Ca²⁺ reacts with HPO₄ ²⁻ to form DCPD on the specimen’s surface (Eq. (5)).

Fig. 1.

GIXRD patterns of the (a) NaOH pre-treated, (b) DCPD-coated, (c) NaOH post-treated DCPD-coated and (d) FHA-coated specimens.

If the pH is locally raised, apatite could be directly agglomerated to the specimen’s surface (Eq. (6)) [29]. Briefly, the presence of apatite phase on DCPD-coated specimen (Fig. 1(b)) is attributed to occasional pH rise. Mg(OH)₂, Mg, distinct apatite accompanied by insignificant DCPD phases are detected on HA-coated specimen (Fig. 1(c)). The presence of DCPD phase proves that the post-treatment partially transformed the DCPD phase to apatite phase (Eq. (7)). Notwithstanding the fact that DCPD is prone to dissolve in human body environment [30], the partial transformation would nurture the process of bone heal with rapid release of Ca²⁺ and HPO₄ ²⁻, while apatite would support the process relatively in medium and long-term $\begin{array}{l} (2) & 2 {OH}^{-} + {Mg}^{+} = Mg {(OH)}_{2}, \\ (3) & 2 H_{2} {PO}_{4}^{-} + 2 e^{-} \to 2 {HPO}_{4}^{2 -} + H_{2} ↑, \\ (4) & 2 {HPO}_{4}^{2 -} + 2 e^{-} \to 2 {PO}_{4}^{3 -} + H_{2} ↑, \\ (5) & {Ca}^{2 +} + {HPO}_{4}^{2 -} + 2 H_{2} O \to {CaHPO}_{4} \cdot 2 H_{2} O (DCPD), \\ (6) & 10 {Ca}^{2 +} + 6 {PO}_{4}^{3 -} + 2 OH \to {Ca}_{10} {({PO}_{4})}_{6} {(OH)}_{2} (HA), \\ (7) & 10 {CaHPO}_{4} \cdot 2 H_{2} O = Ca 10 {({PO}_{4})}_{6} {(OH)}_{2} + 4 H_{3} {PO}_{4} + 18 H_{2} O . \end{array}$

As a result, the HA-coating could have advantages in both short- and long-term. Mg(OH)₂, Mg and distinct apatite phases are detected on FHA-coated specimen (Fig. 1(d)). Remarkably, the crystalline structure of FHA and HA are very similar and the only difference is that F⁻ is substituted with OH⁻. Equations (8)–(10) represent chemical reactions which results in FHA formation. The distinct intensity of apatite peaks in FHA-coating’s GIXRD pattern indicates higher crystallinity in comparison with the HA-coating. Although some previous studies have related the higher crystallinity of FHA to preferential arrangement of apatite crystals, it seems necessary to investigate the phenomenon further. $\begin{array}{l} (8) & 2 F^{-} + {Ca}^{2 +} \to {CaF}_{2}, \\ 5 {CaF}_{2} + (3 - m) H_{2} {PO}_{4}^{-} + m {HPO}_{4}^{2 -} + (1 - x) H_{2} O \to {Ca}_{5} {({PO}_{4})}_{3} {(OH)}_{1 - x} F_{x} \\ (9) & (0 < x < 1, 0 < m < 3), \\ (10) & {Ca}_{5} {({PO}_{4})}_{3} OH + {xF}^{-} + {xH}^{+} \to {Ca}_{5} {({PO}_{4})}_{3} F_{x} {(OH)}_{1 - x} + {xH}_{2} O . \end{array}$

Table 2 provides FWHM and crystallinity degree of apatite crystals embraced in DCPD-, HA- and FHA-coatings. In order to acquire the information, the strong apatite peak at 25.879° was evaluated. According to Table 2, post-treating the DCPD-coating to achieve HA-coating increases the embraced apatite’s crystallinity degree from 47.13% to 68.41%. In addition, substitution of F⁻ with OH⁻ in apatite crystal resulted in enhancement of apatite crystallinity degree from 68.41 to 84.37%. The Higher crystallinity is believed to improve cell growth and decrease the degradation rate of the coatings [31].

Table 2

FWHM and crystallinity degree of apatite in the coatings

Coating	FWHM (β(002))	$X_{c}$ (%)
DCPD coating	0.3083	47.13
HA coating	0.2723	68.41
FHA coating	0.2539	84.37

Table 3 summarizes the results of EDS area analysis, including elemental compositions of calcium, phosphorus and oxygen, as well as Ca/P ratio of the coatings. EDS results reveal that the NaOH post-treatment that transforms DCPD to HA increased the molar ratio of Ca/P of the coating from 1.22 to 1.52, close to theoretical Ca/P ratio of HA (1.67). As a result, dissolution of the coating near the injured bone would nurture the bone with more suitable chemical composition. Nonetheless, Ca/P ratio of the FHA-coating was measured 1.47. The lower Ca/P ratio of the FHA- and HA-coatings in comparison with their theoretically calculated ratio (1.67) implies calcium deficiency in the coatings. As well, non-stoichiometric structures of apatite, responsible for calcium deficiency, are reported common when F⁻ is substituted with OH⁻. Despite the calcium deficiency of the coatings, it is reported that even calcium deficient apatite with Ca/P ratio between 1.33 and 1.55 are able to assist bone heal process [32]. In addition, fluorine was detected in FHA-coatings, proving its presence in the coating structure. The F/P ratio of FHA was about 0.18. It is worth noting that the high antibacterial activity of FHA coating without any harmful effect and positive stimulating effect on cell proliferation were studied [33]. Furthermore, the optimized concentration of F/P ratio of FHA has been reported to be less than 0.317.

Table 3

EDS results as well as experimental and theoretical comparison of Ca/P ratio of the investigated coatings

Coatings	Ca (at.%)	P (at.%)	O (at.%)	Mg (at.%)	F (at.%)	Ca/P (Molar ratio)	Theoretical Ca/P (Molar ratio)
DCPD	14.01 ± 0.09	11.42 ± 0.04	58.78 ± 0.15	15.80 ± 0.02	–	1.22	1
HA	13.34 ± 0.06	8.77 ± 0.08	68.85 ± 0.21	9.04 ± 0.13	–	1.52	1.67
FHA	14.22 ± 0.07	9.64 ± 0.11	68.13 ± 0.18	6.31 ± 0.01	1.76 ± 0.01	1.47	1.67

Nevertheless, a considerable amount of magnesium was also detected, most probably because the coatings are porous and low thickness compared to the EDS detection depth (micrometer scale). The decrease in the amount of magnesium detected in the FHA coating implies that the resulted coating is more compact in comparison with HA coating and definitely more compact than DCPD coating. Finally, fluorine was detected in FHA-coatings proving its presence in the coating structure.

Figure 2 illustrates the ATR-FTIR spectra of the NaOH pre-treated and coated specimens. The broad bands around 3250 cm⁻¹ are attributed to the adsorbed H₂O [34]. For the spectra of the coated specimens (Fig. 2(b)–(d)), the multiple bands around 600 and 1000 cm⁻¹ are characterized as PO₄. These split bands specify well-crystallized HA. In addition, the spectra show the presence of P–OH band at around 870 cm⁻¹ which is ascribed to HPO₄ ²⁻ in the coatings, revealing other Ca–P compositions in the FHA coating that were not detected by GIXRD. In addition, there are bands of OH⁻ at about 1300 and 3500 in the spectra of DCPD- and FHA-coated specimens (Fig. 2(b) and (d)). Comparison of Fig. 2(b) and (d) also clearly demonstrates decreasing in the absorbance of the OH⁻ band at around 1200 and 3500 cm⁻¹ due to substitution of OH⁻ by the F⁻. The strong bands at around 1400 cm⁻¹ in Fig. 4(a) and (c) indicate significant absorbance of OH⁻. In addition, in Fig. 2(c), the strong peak around 3700 cm⁻¹ refers to OH⁻.

Fig. 2.

ATR-FTIR spectra of the (a) NaOH pre-treated, (b) DCPD-coated, (c) HA-coated, and (d) FHA-coated specimens.

3.2. Morphology characterization

Figure 3 illustrates photographs of the surface of the NaOH pre-treated and coated specimens. Using arithmetic means, the thickness of the DCPD, HA and FHA coatings were measured as 17.6, 13.8 and 23.4 µm, respectively. Figure 4 illustrates the cross-sectional optical micrographs of the coated specimens. Small changes in the thickness of the coatings can be observed. Noticeably, due to being dissolved in NaOH solution, the thickness of the HA coating is reduced in comparison with DCPD coating.

Fig. 3.

Photographs of the NaOH pre-treated and coated specimens.

Fig. 4.

Cross-sectional optical micrographs of the (a) DCPD, (b) HA and (c) FHA-coated specimens.

Morphology of the coated specimens is illustrated in Fig. 5. It can be seen from Fig. 5(a) that flake-like DCPD coating was formed on the magnesium surface without delamination. The surface morphology of HA-coated specimen is shown in Fig. 5(b). It is notable that the HA and DCPD coatings’ morphologies are relatively similar, flake-like, while HA coating has a packed flake-like morphology.

Fig. 5.

SEM images showing morphology of the (a) DCPD, (b) HA and (c) FHA-coated specimens; (d) high magnification of FHA.

The compactness of HA coating stems from flake-like morphology of DCPD that was dissolved to some extent and as a result HA was regenerated on the specimen’s surface. The effect of NaOH post-treatment can be described by dissolving some of the DCPD coating into NaOH solution followed by releasing Ca²⁺ and HPO₄ ²⁻ into the solution leading to localized increase in Ca²⁺ and PO₄ ³⁻. This increase in the concentration of ionic species results in the nucleation and growth of HA [8]. The DCPD phase transformation into the HA phase is explained according to Eq. (7). Figure 5(c) represents the morphology of FHA coating. Figure 5(c) clarifies coexistence of 2 different layers in FHA coating. The first layer consists of closely paced and directionally arranged columnar crystals (indicated as interlayer). The second layer comprises irregularly nucleated crystals that are accumulated over the interlayer. The FHA crystals morphology is described in high magnification by Fig. 5(d). Conspicuously, the morphology of FHA has been reported as a suitable ground for the precipitation of Ca–P during immersion in SBF due to the significant increase in the surface area leading to development of a coating which decreases the degradation of the FHA coating [35]. The external layer of FHA coating is compact with a small crystal size compared to the HA coating. This is in strong agreement with the results of other researchers [36] who ascribed the formation of small size crystals to the small ion size of F⁻ (1.32 Å), compared to OH⁻ (1.68 Å), which can decrease the crystal size, and consequently produce denser structure.

3.3. Topography characterization

Surface topography of the uncoated, NaOH pre-treated and coated specimens were investigated by AFM and are shown in Fig. 6. Scratches originated from the grinding process can be seen on the surface topography of the uncoated specimen (Fig. 6(a)). The surface topography of NaOH pre-treated specimen (Fig. 6(b)) is different from the surface topography of the uncoated specimen. The coatings altered the specimens’ surface topography fundamentally (Fig. 6(c)–(e)). The RMS value of the ground specimen was 238 nm, while that of NaOH pre-treated specimen was measured 292 nm. The RMS value of the specimens coated with DCPD, HA and FHA increased to 384, 437 and 502 nm, respectively. The DCPD-coated specimens showed moderately porous and multi-directional flake-like crystallization which resulted in smooth surface.

Fig. 6.

AFM topography of the (a) uncoated, (b) NaOH pre-treated, (c) DCPD-coated, (d) HA-coated and (e) FHA-coated specimens.

Conversely, HA-coated specimen underwent NaOH post-treatment which dissolved DCPD coating to some extent and resulted in a relatively rough surface. Thus, RMS value of HA-coated specimen increased in comparison with that of DCPD-coated specimen. On the other hand, the FHA coating was neither porous enough nor multi-directionally crystallized. As a result, the RMS value was increased significantly. It is worth noting that surface roughness is one of the factors affecting implants’ biocompatibility. The higher surface roughness of an implant brings about the more blood coagulation and it is favoured for implantation in hard tissue [37]. Rough surfaces are also favoured in orthopaedic applications because of the enhanced adhesion between implants and osteoblast cells [38].

3.4. Wettability characterization

The first contact of implant with physiological environment of human body after placement in its location is with physiological fluids. Wetting by physiological fluids is important as it would control adsorption of proteins which introduce cell attachment to implants’ surface. The results showed acceptable accuracy, less than 10% error. The uncoated specimen showed the highest contact angle ( $122.3 \pm 8.6 °$ ), and only through NaOH pre-treatment, the specimen’s contact angle decreased to $87.3 \pm 7.9 °$ (Fig. 7). The contact angle of the DCPD-coated and HA-coated specimens was $64.7 \pm 3.1 °$ and $44.1 \pm 1.6 °$ , respectively. For untreated, NaOH pre-treated, DCPD and HA-coated specimens, the higher the RMS value, the lower the contact angle. Conversely, although, the FHA-coated specimen had the highest RMS value, the contact angle of it was $49.8 \pm 1.5 °$ . Contact angle concept could be used to evaluate wettability phenomenon through interfacial tensions of specimen-drop ( $γ_{sl}$ ), specimen-air ( $γ_{sv}$ ) and drop-air ( $γ_{lv}$ ). Young’s Eq. (11) correlates the interfacial tensions and contact angle (θ). $\begin{matrix} (11) & cos θ = γ_{sv} - γ_{sl} / γ_{lv} . \end{matrix}$

The γ _lv is a constant positive value because laboratory environment is controlled and the same liquid, water, is used for every test. In addition, if $θ < 90 °$ , the specimen’s surface is considered as wettable or hydrophilic. If $0 ⩽ cos θ ⩽ 1$ , according to Eq. (11), wettability could be determined merely by the difference between γ _sv and γ _sl, known as solid surface tension. Chemical composition of specimens’ surface, microstructural topography and surface charge are the most prominent factors to vary the solid surface tension and wettability. In regard to the higher contact angle of FHA than HA in this study, previous researchers attributed the higher contact angle of FHA-coated specimen to replacement of OH⁻ groups by F⁻, and increasing the negative charge on the surface [39]. It is believed that the lower the contact angle, the more the quantity of adhered osteoblasts as well as cell activity [40]. Therefore, the HA and FHA coatings are able to provide an improved mechanism of osseointegration.

Fig. 7.

Illustration of contact angles corresponding to the median of the measurements; (a) magnesium, (b) pre-treated, (c) DCPD-coated, (d) HA-coated, (e) FHA-coated substrate.

3.5. Corrosion characterization

Figure 8 illustrates the electrochemical polarization curves of the specimens obtained through the potentiodynamic polarization test performed in Kokubo SBF. Although, the tests were repeated three times for each surface condition to eliminate any deviation, the results showed acceptable consistency and no data elimination was necessary. Polarization curves of all the specimens show comparable electrochemical corrosion behaviors. Shifting in the polarization curves of the coated specimens to higher corrosion potential (E_corr) as well as lower corrosion current density (I_corr), are clearly demonstrated. It can also be observed that shifting the polarization curve towards an improved corrosion resistance position is the most for the FHA-coated specimen.

Fig. 8.

Electrochemical corrosion behavior of the uncoated, pre-treated and coated specimens in Kokubo SBF.

The shift for that of the HA-coated specimen is stood in the second position. While the DCPD-coated specimen experienced lower shifting compared to the other coated specimens. In other words, although the Mg(OH)₂ layer formed on the surface of magnesium after NaOH pre-treatment seems to impede corrosion, it could not sufficiently protect the specimen. The improvement in electrochemical corrosion behavior of the HA-coated specimen, compared to the DCPD-coated specimen, is attributed to relative compactness and uniformity of the coating (Fig. 5(b)). The relatively compact and uniform HA coating hinders ionic diffusion and electrochemical reactions between the media and specimen and results in higher corrosion resistance. Furthermore, the pronounced progress in corrosion resistance achieved from FHA-coated specimen, in comparison with DCPD and HA-coated specimens, is ascribed to its higher dissolution resistance. Dissolution resistance of FHA-coating stems from presence of F⁻ in the coating structure [41]. For a more explicit comparison, Table 4 summarizes the arithmetic average of E_corr (in terms of V vs. SCE; V_SCE) and I_corr (in terms of A/cm²) extrapolated from the polarization curves. It can be seen that E_corr of the uncoated magnesium increased from −2.02 V_SCE to −1.44 V_SCE for FHA-coated specimen. In addition, I_corr of the uncoated magnesium decreased from $3.91 \times 10^{- 5} {A/cm}^{2}$ to $8.72 \times 10^{- 6} {A/cm}^{2}$ for the DCPD-coated and to the $1.46 \times 10^{- 6} {A/cm}^{2}$ for FHA-coated specimen. These results indicate that magnesium corrosion remarkably decreases, about 4.5 times, when coated by DCPD, and even more remarkably decreases, about 27 times, when coated by FHA.

Table 4

I_corr and E_corr of the specimens extracted from the polarization curves

Electrochemical parameter	Mg	NaOH pre-treated Mg	Coating material

			DCPD	HA	FHA
I_corr (A/cm²)	$3.91 \times 10^{- 5}$	$3.31 \times 10^{- 5}$	$8.72 \times 10^{- 6}$	$4.60 \times 10^{- 6}$	$1.46 \times 10^{- 6}$
E_corr (V_SCE)	−2.0272	−1.8655	−1.63394	−1.53395	−1.44156
Corrosion rate (mmpy)	2.19	1.85	0.48	0.25	0.08

3.6. The pH evaluation

The results of pH measurement are shown in Fig. 9. It can be seen that pH value of all specimens corresponding SBF increased after 14 days of immersion. Significantly, during the first day of immersion, pH value of the coated specimens corresponding SBF rapidly increased, and proved dissolution of the coatings and release of OH⁻. However, extension of immersion time decreased the slope of pH elevation which is most probably due to the penetration of SBF into coating–specimen interface. In other words, presence of SBF in coating–specimen interface brings about pH elevation – adjacent to the specimen’s surface – disturbing anodic reactions (Eq. (12)). In general, during magnesium corrosion, cathodic reaction, according to Eq. (13), generates OH⁻ leading to elevation of pH and acceleration of hydrogen (H₂) evolution [42].

Fig. 9.

The pH value of Kokubo SBF plotted as a function of immersion time for the pure, NaOH pre-treated as well as coated magnesium.

It should be noted that Mg²⁺ reacts with OH⁻ and forms Mg(OH)₂ layer (Eq. (14)) on the surface of specimens. While Mg(OH)₂ layer could protect specimens’ surface, the aggressive environment of human body – resembled by SBF – attacks the protective layer and renews specimens’ corrosion. $\begin{array}{l} (12) & Mg = {Mg}^{2 +} + 2 e, \\ (13) & 2 H_{2} O + 2 e = H_{2} + 2 {OH}^{-}, \\ (14) & {Mg}^{2 +} + 2 {OH}^{-} = Mg {(OH)}_{2} . \end{array}$

SBF contains a considerable amount of chloride (Cl⁻) which attacks and converts Mg(OH)₂ to MgCl₂ (Eq. (15)). Conversion of Mg(OH)₂ to MgCl₂ releases OH⁻ which increases pH of SBF. In addition, MgCl₂ dissolves easily in SBF, according to Eq. (16), and releases Cl⁻. The resulted Cl⁻ from MgCl₂ dissolution continuously attacks Mg(OH)₂, the protective layer on specimens surface, and by weakening Mg(OH)₂ layer renews the specimens’ corrosion. The process of weakening Mg(OH)₂ layer and renewing specimens’ corrosion adds a considerable amount of OH⁻ and Mg²⁺ to SBF, especially adjacent to specimens’ surface. Accordingly, pH of SBF increases exponentially during the first days of immersion experiment. Increasing pH and Mg²⁺ concentration results in the formation of Ca–P compositions through reactions between Ca²⁺, HPO₄ ²⁻ or PO₄ ³⁻, OH⁻ and Mg²⁺ ions. Formation and recrystallization of Ca–P on uncoated and coated specimens, respectively, retards the aggressive corrosion during the first days. $\begin{array}{l} (15) & Mg {(OH)}_{2} + {Cl}^{-} = {MgCl}_{2} + {OH}^{-}, \\ (16) & {MgCl}_{2} = {Mg}^{2 +} + {Cl}^{-} . \end{array}$

Among coated specimens, the DCPD-coated specimen showed the highest pH of SBF in immersion test and is the most susceptible to corrosion in SBF in comparison with HA and FHA-coated specimens, which is in consistence with the results of electrochemical corrosion test. Porosities in the DCPD coating most probably let SBF penetrate to specimen-coating interface, thereby causing higher pH and corrosion. However, extension of pH evaluation test results in excessive DCPD dissolution in SBF and pH elevation. Consequently, Ca–P compositions precipitate on coating defects and porosities and decelerate pH elevation and corrosion. The newly precipitated Ca–P compositions not only limit exposure of specimens to SBF but also stabilise the DCPD coating and eventually cause improvement of DCPD-coated specimen corrosion resistance. HA-coated specimen showed considerably higher corrosion resistance in comparison with the uncoated specimen. In this regard, HA revealed a more protective coating layer compared with DCPD. The comparably lower pH of SBF corresponding to HA stems from the compactness and uniformity of HA coating which prevents exposure of the specimen to SBF. In addition, HA structure is more stable and less dissolvable in SBF thus the HA-coated specimen is protected by a stabilized protective layer. The FHA-coated specimen showed the least pH elevation among the coated specimens and revealed pronounced corrosion resistance over the uncoated specimen. With the lowest solubility in SBF and due to the dense, stable and close packed structure of FHA, this material resists against dissolution and prevents pH elevation. As well, the integrated FHA coating protects magnesium and denies significant SBF exposure. It is worth noting that the higher stability of FHA compared to HA layer is attributed to the crystal structure of FHA and faster dissolution of Ca²⁺ for HA with increased incubation time.

4. Conclusions

In this study, different types of Ca–P phases were coated on NaOH pre-treated magnesium by electrodeposition technique. The coatings showed significant improvements on the corrosion resistance, roughness and wettability of magnesium. DCPD and FHA were found deposited on magnesium specimens; therefore, NaOH post-treatment was necessary on the DCPD-coated specimens to recrystallize HA. The results suggest that the coatings radically alter surface topography of the specimens and considerably increase the corresponding roughness. FHA coating showed higher surface roughness than those of DCPD and HA coatings. In addition, all of the coatings changed specimens’ wettability from a hydrophobic state to a defined hydrophilic state. Although surface roughness impacts directly on wettability, HA-coated specimen showed the highest hydrophilicity, due to the F⁻ substitution with OH⁻. In regard to corrosion performance of the coated specimens, it was revealed that coatings enhanced corrosion resistance in both electrochemical and immersion evaluations. The highest pronounced improvement in corrosion resistance, about 27 times, was observed in FHA-coated magnesium. In addition, it was observed that an increase in crystallinity of the coatings (FHA and HA compared to DCPD) improved corrosion resistance of the specimens by decreasing the coatings solution in SBF.

Footnotes

Acknowledgements

The authors would like to acknowledge the Ministry of Higher Education of Malaysia for the financial support (vote number 04H18) and faculty of mechanical engineering of the Universiti Teknologi Malaysia for providing research facilities, and Shahid Rajaee Teacher Training University for the helpful collaboration.

Conflict of interest

The authors have no conflict of interest to report.

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	Ion concentration (mmol/L)

	Na⁺	K⁺	Ca²⁺	Mg²⁺	HCO³⁻	Cl⁻	HPO₄ ²⁻	SO₄ ²⁻
Kokubo SBF	142.0	5.0	2.5	1.5	4.2	147.8	1.0	0.5
Plasma solution	142.0	5.0	2.5	1.5	27.5	103.0	1.0	0.5

Topography,wetting,and corrosion responses of electrodeposited hydroxyapatite and fluoridated hydroxyapatite on magnesium

Abstract

Keywords

1. Introduction

2. Experimental methods and materials

2.1. Specimen and coating electrolyte preparation

2.2. Coating deposition

2.3. Coating characterization

2.4. Electrochemical corrosion tests

Table 1 Chemical composition of Kokubo SBF and human blood plasma [28] Ion concentration (mmol/L) Na+ K+ Ca2+ Mg2+ HCO3− Cl− HPO4 2− SO4 2− Kokubo SBF 142.0 5.0 2.5 1.5 4.2 147.8 1.0 0.5 Plasma solution 142.0 5.0 2.5 1.5 27.5 103.0 1.0 0.5

3.1. Crystalline and chemical characterization

Footnotes

Acknowledgements

Conflict of interest

References

Table 1
Chemical composition of Kokubo SBF and human blood plasma [28]

Ion concentration (mmol/L)

Na⁺ K⁺ Ca²⁺ Mg²⁺ HCO³⁻ Cl⁻ HPO₄ ²⁻ SO₄ ²⁻

Kokubo SBF 142.0 5.0 2.5 1.5 4.2 147.8 1.0 0.5

Plasma solution 142.0 5.0 2.5 1.5 27.5 103.0 1.0 0.5