Development of HA-CNTs composite coating on AZ31 Magnesium alloy by cathodic electrodeposition. Part 2: Electrochemical and in-vitro behavior

Ceramics International 45 (2019) 11186–11194

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Development of HA-CNTs composite coating on AZ31 Magnesium alloy by cathodic electrodeposition. Part 2: Electrochemical and in-vitro behavior

T

D. Khazeni, M. Saremi , R. Soltani ⁎

School of Metallurgy and Materials Engineering, College of Engineering, University of Tehran, 11155-4563 Tehran, Iran

ARTICLE INFO

ABSTRACT

Keywords: Magnesium alloy Electrodeposition Hydroxyapatite Carbon nanotube Corrosion

The carbon nano-tubes (CNTs) reinforced hydroxyapatite (HA), with various functionalized CNTs concentration ranging from 0 to 1.5 wt%, were deposited on AZ31 magnesium alloy by direct and pulse cathodic electrodeposition methods. The corrosion resistance of the coatings was tested in simulated body fluid (SBF) using different electrochemical methods such as open circuit potential, polarization and electrochemical impedance spectroscopy. The in-vitro behavior, changes in solution pH as well as the amount of evolved hydrogen of these coatings were also evaluated during five days immersion in SBF. The results indicated that the pulse deposited HA having 1% CNTs coating was the optimum condition which decreased the corrosion current density of AZ31 magnesium alloy from 44.25 µA/cm2 to 0.72 µA/cm2. Moreover, it stabilized the alkalization behavior of AZ31 alloy and caused a tenfold decrease in the amount of hydrogen generation in SBF. Additionally, the formation of new hydroxyapatite layer on the surface of the pre-exist coatings after five days immersion in SBF was confirmed by SEM characterization.

1. Introduction Classical biomedical implants, such as stainless steel and titanium alloys, play a crucial role in biomedical application especially in repairing the damaged human bone tissue [1–5]. However, there are concerns about toxicity and the released ions from such implants in long term use [6,7]. Additionally a second surgery for removing the implant after bone curing is also another problem which necessitates the research for biodegradable implant [8]. Recently, various types of biodegradable polymers are used as temporary implants but they have problems of inflammation and lack of mechanical strength where used in load-bearing parts [9,10]. Whereas Mg and its alloys reveal a promising candidate as a biodegradable implant thanks to their ability for degradation in addition to their appropriate mechanical properties as well as good biocompatibility [11–13]. However, magnesium and its alloys are vulnerable to rapid corrosion in a human body fluid or blood plasma. It is reported that these alloys have a strong tendency for non-homogeneous and localized corrosion in the environments containing chloride which results in losing mechanical properties before the healing of the bone tissue [14,15]. Another major issue concerning magnesium alloys is the formation of hydrogen gas during the corrosion process [16,17]. Generally, Mg and its alloys exhibit a negative open circuit potential in

⁎

aqueous solutions that are reported to be between −1.2 and −1.7 V in SBF [18,19]. This potential is low enough at all values of solution pH to provide the thermodynamic conditions for dissociation of water and subsequently hydrogen gas formation. In this condition, the balloon effect may occur due to high hydrogen liberation along with the change in solution pH to more alkaline in the vicinity of corroding regions [20]. These reactions would hinder the growth and healing process of bone tissue which is harmful in medical applications [18]. There are many strategies to improve magnesium corrosion in human body's environment. Application of coatings on Mg and its alloys is a suitable method to overcome corrosion [21]. Recently, many researchers draw their attention to Hydroxyapatite [HA, Ca10(PO4)6(OH)2] coating as a biocompatible composition which could also improve the corrosion resistance of magnesium alloys [22–24]. However, the mechanical strength of HA is too low to be employed in load-bearing applications [25]. To this end, the addition of reinforcing materials like CNTs with its excellent mechanical and physical characteristics is considered as a new approach to improve HA mechanical properties [26,27]. The biocompatibility of the implant intended for orthopedic is a significant parameter. Both Mg alloys and hydroxyapatite are proven to be biocompatible, but the biocompatibility of CNT is still under debate [28,29]. There are three main reasons regarding the application of

Corresponding author. E-mail address: [email protected] (M. Saremi).

https://doi.org/10.1016/j.ceramint.2019.01.105 Received 15 November 2018; Received in revised form 4 January 2019; Accepted 14 January 2019 Available online 22 January 2019 0272-8842/ © 2019 Elsevier Ltd and Techna Group S.r.l. All rights reserved.

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HA–CNTs composite in orthopedic devices. (i) The CNTs possess cytotoxic effect while they are directly suspended in a fluid medium. However, they appeared to be nontoxic if immobilized to a matrix [30]. (ii) There are many reports of CNTs having a positive response to bones and bone cells in open literature [31–33]. Gopi et al. developed HACNTs composite coatings having 0.1–2 wt% CNTs on Ti substrate by electrodeposition method. They proved that CNTs plays a significant role in the enhancement of cell viability in HA–CNTs composite coating [34]. Pei et al. claimed that HA-CNTs composite coatings induced better cell proliferation, cell viability and ALP activity in comparison to pure HA coating [35]. (iii) If the release of CNTs occurs in the blood due to the wear of the composite surface, they are entirely biodegraded by neutrophils and macrophages [28] and safely and rapidly purge from the body by renal excretion, without causing any toxic effect in the human body [36,37]. In this regards, adding CNTs to HA composite should not have a negative impact on the biocompatibility. Several researchers have worked on the corrosion behavior of magnesium alloys coated by different types of calcium phosphate coatings [20,38]. Meng and his coworkers [39], showed that the fluorine-doped hydroxyapatite coating obtained by electrodeposition method improved the corrosion resistance of the Mg–Zn–Ca alloy. Wu et al. [40], prepared calcium phosphate coatings on AZ91 substrate by pulse electrodeposition method at different duty cycles and reported that the coating obtained by duty cycle of 0.10 showed the best corrosion resistance. Shangguan and his coworkers [41] fabricated a brushite coating on the Mg-Sr alloy by electrodeposition method and surprisingly reported that the corrosion rate of the coated alloy was increased compared to the bare Mg-Sr alloy. They explained that the formation of the inner corrosion layer beneath the brushite layer during the electrodeposition process was the reason for such behavior. It was concluded that the magnesium alloys with poor corrosion resistance are not suitable to be coated by electrodeposition methods. To the best of our knowledge, there is no comprehensive report on microstructural and corrosion behavior of electrodeposited HA–CNTs on AZ31 Magnesium alloy. The objective of these two parts series papers is to develop CNTs reinforced HA composite coatings on AZ31 magnesium alloy using direct and pulse electrodeposition methods. In the second part, the effect of CNTs and its amount in the coating composition on the corrosion resistance and in-vitro behavior of the specimens are emphasized. 2. Experimental procedure 2.1. Development of HA and HA-CNTs composite coating on AZ31 alloy The AZ31 magnesium samples were cut in 2 cm × 2 cm workpieces, polished by silicon carbide papers (180–1200 grits), cleaned in an alcohol-acetone mixture, rinsed with distilled water and dried in air. An area of 1 cm2 was subjected to the tests and the rest was covered by a resistant locker. The electrolyte for hydroxyapatite deposition consists of 0.042 mol/ L Ca(NO3), 0.025 mol/L NH4H2PO4 and 10 ml/L H2O2 and the pH of the electrolyte was adjusted to 4.7 by dilute HNO3 and NH3. All of the experimental processes were done at ambient temperature. For fabricating composite coating on the AZ31 substrate, CNTs with carboxyl functionalized group were gradually added to the coating bath at different concentrations of 0.25, 0.5, 1 and 1.5 wt%. In order to obtain a more homogenous distribution of CNTs, prior to any electrodeposition process, the electrolyte was subjected to an ultrasonic treatment for 30 min. Electrodeposition processes were performed in a cell using a regular three-electrode configuration in which AZ31 magnesium alloy was employed as the cathode, a platinum electrode as the anode and a saturated calomel electrode (SCE) as the reference electrode and the solution was stirred at 400 rpm. The voltage in direct mode was adjusted to −3V, while for pulse mode the condition was set to an optimized

Table 1 The samples label according to coating condition. Code name

Testing condition

Voltage

D-CP D-HA D-HA-C0.25 D-HA-C0.5 D-HA-C1 D-HA-C1.5 P-CP P-HA P-HA-C0.25 P-HA-C0.5 P-HA-C1 P-HA-C1.5

Calcium phosphate coating prior to alkalization Pure Hydroxyapatite coating Hydroxyapatite coating + 0.25 wt% CNTs Hydroxyapatite coating + 0.5 wt% CNTs Hydroxyapatite coating + 1 wt% CNTs Hydroxyapatite coating + 1.5 wt% CNTs Calcium phosphate coating prior to alkalization Pure Hydroxyapatite coating Hydroxyapatite coating + 0.25 wt% CNTs Hydroxyapatite coating + 0.5 wt% CNTs Hydroxyapatite coating + 1 wt% CNTs Hydroxyapatite coating + 1.5 wt% CNTs

Direct Direct Direct Direct Direct Direct Pulse Pulse Pulse Pulse Pulse Pulse

−3V with a duty cycles of 0.2 at room temperature. Finally, the samples subjected to alkaline treatment for two hours in one molar sodium hydroxide solution at 80 °C. 2.2. Sample Identification As is shown in Table 1 the coated samples are abbreviated in accordance with their testing condition. Meanwhile, the term “AZ31” referred to the uncoated AZ31 magnesium, as-received sample, in the article. 2.3. Electrochemical measurements The electrochemical measurements such as open circuit potential (OCP), potentiodynamic polarization and electrochemical impedance spectroscopic (EIS) tests were carried out to evaluate the corrosion behavior of the coatings. In this regard, a three-electrode corrosion cell was used for electrochemical measurements as well. The counter electrode was made of platinum and the saturated silver/silver chloride electrode was employed as the reference electrode. The exposed area of the working electrode (coated AZ31 alloy) to the solution was 1 cm2. All the electrochemical tests were conducted in the simulated body fluid (SBF) and the temperature was 37 ± 1 °C. The composition of SBF solution is shown in Table 2. The OCP measurement was performed during 15 min immersion of the samples in SBF. The polarization test was conducted at a constant scan rate of 0.5 mV/s by means of an EG&G potentiostat, model 273 A. Also, the electrochemical impedance spectroscopy (EIS, Solarton1260) in the frequency range between 100 kHz and 10 mHz with an amplitude of ± 5 mV peak-to-peak, was employed. 2.4. pH and hydrogen evolution measurement A group of specimens was immersed in SBF for 5 days to evaluate the changes in solution pH and H2 evolution. The pH of the solution was measured with a portable pH meter and the hydrogen generation was assessed as per the method suggested in Ref [19]. 2.5. Microstructure characterization The morphology of the coated samples after 5 days immersion in SBF solution and its elemental composition were characterized by SEM Table 2 Chemical composition of SBF solution (all of the contests are mM except for pH).

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Na+

K+

Mg2+

Ca2+

Cl-

HCO32-

HPO42-

SO42-

pH

142

5

1.5

2.5

148.8

4.2

1

0.5

7.4

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Potential (v vs. SCE)

-1.25

(a)

-1.35 -1.45

AZ31 (1) D-CP (2)

4 3 2

-1.55

D-HA (3) D-HA-C0.25 (4)

1

-1.65

D-HA-C0.5 (5) D-HA-C1 (6)

-1.75

D-HA-C1.5 (7)

-1.85

0

-1.25 Potential (v vs. SCE)

6

7 5

100 200 300 400 500 600 700 800 900 Time (s)

(b)

-1.35

6

-1.45

2

7 5 4 3

Fig. 3. Electrochemical parameters extracted from polarization tests data for uncoated AZ31 magnesuim alloy and different coatings conditions.

AZ31 (1) P-CP (2) P-HA (3)

-1.55

(Tescan Vegal) with Energy Dispersive Spectroscopy (EDS).

P-HA-C0.25 (4)

1

-1.65

P-HA-C0.5 (5)

-1.75

P-HA-C1.5 (7)

P-HA-C1 (6)

-1.85

0

3. Results and discussion 3.1. Electrochemical studies

100 200 300 400 500 600 700 800 900 Time (s)

Fig. 1. OCP time plots obtained in SBF solution for uncoated AZ31 magnesuim alloy and different coatings conditions: (a) Direct Voltage, (b) Pulse Voltage.

-0.8

(a)

E (V vs. Ag/Agcl)

-1

7 4

6

-1.2

AZ31 (1)

2

3

5

D-CP (2)

1

D-HA (3)

-1.4

D-HA-C0.25 (4)

Anodic reaction: Mg → Mg2+ + 2e-

D-HA-C0.5 (5)

-1.6 -1.8 -2

-7

-6

-5

-4

-3

-2

-1

0

-

(b)

Cathodic reaction: 2H2O + 2e → 2OH + H2

(2)

D-HA-C1.5 (7)

Over reaction: Mg + 2H2O → H2 + Mg(OH)2

(3)

Furthermore, Mg(OH)2 is formed by the conversion of oxide film (MgO) on the surface of bare AZ31 through the following reaction (Eq. (4)) [17]:

1

MgO + H2O → Mg(OH)2

6

3

E (V vs. Ag/Agcl)

-1 7

4

2 AZ31 (1)

5

-1.2

P-CP (2)

1

P-HA (3)

-1.4

P-HA-C0.25 (4) P-HA-C0.5 (5)

-1.6

P-HA-C1 (6) P-HA-C1.5 (7)

-1.8 -2

-7

-6

-5

-4

-3

-2

-1

0

(1) -

D-HA-C1 (6)

Log i (A/cm2) -0.8

3.1.1. Open-circuit potential (OCP) measurements OCP measurement is a simple and notable method to evaluate the stability of metals and coatings in corrosive media. Figs. 1a and 1b show the potential-time variation of different samples in SBF solution during 15 min immersion at 37 ± 1 °C. The open-circuit potential of the uncoated AZ31 magnesium alloy is more active than that of the coated samples. At the early stage of immersion, AZ31 magnesium alloy shows an initial increasing trend to more positive potentials. This is probably due to the formation of some corrosion products on the surface. The following reactions show the corrosion sequence of magnesium alloy in aqueous solution (Eqs. (1)–(3)) [17]:

1

Log i (A/cm2) Fig. 2. Potentiodynamic polarization curves in SBF solution for uncoated AZ31 magnesuim alloy and different coatings conditions: (a) Direct Voltage, (b) Pulse Voltage.

(4)

It is seen that, however, the magnesium oxide/hydroxide layer formed on substrate reduce the surface activity in further stages of immersion, the open-circuit potential of AZ31 is still far less than the coated samples. This happen because the porous structure of the corrosion products cannot act as a compact oxide layer leading to more corrosion of AZ31 alloy [40]. Some fluctuations are seen in OCP of the coated samples which may be due to the occurrence of unstable localized corrosions (decreasing potential) followed by the formation of corrosion products (increasing potential) under the defects of the coatings. The most fluctuation occurred in D-CP sample, which indicates the poor stability of DCPD phase in SBF solution, while P-CP sample shows less fluctuation probably due to the presence of HA phase along with DCPD phase in this coating. It is in agreement with the fact that HA phase is more stable in the solution with pH around 7 than DCPD phase [42]. Moreover, by comparing Figs. 1a and 1b, it can be concluded that the samples coated by pulse voltage electrodeposition show less fluctuation and have more positive potential than that of coated by the direct method. Based on the microstructural characterization described in the previous part of

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Fig. 4. SEM images of composite coating (a) D-HA, (b) P-HA, (c) D-HA-C1 and (d) P-HA-C1 samples after potentiodynamic polarization test.

this research, the major reason is the morphology of the pulse electrodeposited coatings which is denser and more uniform compared to the direct method. Furthermore, during the pulse deposition method, H2 formation is declined leading to the formation of more homogenous and defect-free coatings. Regarding CNTs effect, the presence of carbon nanotubes in the coatings enhances the potential to the more positive values showing that HA-CNTs coatings provide a superior barrier against the corrosive solution than pure HA coatings. This may be described by the bridging mechanism of CNTs between the HA crystals which makes the coating structure more compact and reduces the access of corrosive solution to the base material. By increasing the CNTs amount up to 1% in the composite coatings, the potentials of the samples increase in which PHA-C1 sample shows a noble OCP value (−1.32 V) as well as a most stable trend during the immersion in SBF solution. While by adding more CNTs, 1.5%, the potential of the coatings experienced a decrease in both direct and pulsed voltage method. This can be related to the non-uniform coating morphology of D-HA-C1.5 and P-HA-C1.5 samples due to the agglomeration of CNTs. 3.1.2. Potentiodynamic polarization tests Figs. 2a and 2b show Tafel polarization curves of AZ31 magnesium alloy and the coated samples obtained from direct and pulse voltage electrodeposition after 15 min immersion in SBF solution at 37 ± 1 °C. The variation of the polarization parameters extracted from the Tafel polarization curves are shown in Fig. 3 for each sample. First of all, the Ecorr of all the coated samples are shifted to more positive potentials and a significant reduction is occurred in their icorr compared to AZ31 magnesium alloy. Moreover, the pulse deposited coatings show better

corrosion behavior compared to direct ones. Additionally, the presence of CNTs reduces the dynamic and kinetic tendency for corrosion which is concluded from Ecorr and Icorr, respectively. To be more precise, the crystallinity of the coatings can be considered as another factor affecting the corrosion resistance of the samples in addition to the aforementioned parameters explained in the OCP section. As it was discussed in part 1 of this research, SEM images show that the functionalized CNTs increase the nucleation sites of hydroxyapatite during the electrodeposition process and alkaline treatment applied after deposition. Increasing the hydroxyapatite nuclei may reduce and fill the passageways in the composite coatings where the SBF solution can get into the substrate. Moreover, the XRD results confirm that by increasing the amount of CNTs in the HA-CNTs composite coatings, the crystallinity of the coatings is increased. Some researchers reported that higher crystallinity leads to greater stability of hydroxyapatite in simulated body fluids [43]. In the present research it is seen that the crystallinity is increased in the coatings containing 1.5% CNTs, however, the corrosion resistance is decreased due to the agglomeration of CNTs and less homogeneity of the coatings in both direct and pulse electrodeposition methods. The most protective coating against corrosion was the pulse deposited hydroxyapatite having 1% CNTs. The icorr and Ecorr of P-HA-C1 sample are 0.72 µA/cm2 and −1.31 V vs. Ag/AgCl, whereas the icorr and Ecorr values of the uncoated AZ31 magnesium alloy are found to be 44.25 µA/cm2 and −1.57 V vs. Ag/AgCl, respectively. Fig. 4 shows the SEM images of the coated samples after Tafel polarization test, in which the pits are shown by white arrows. It is seen that more local corrosion occurred in the coatings obtained from direct voltage deposition in comparison with the pulse deposited coatings.

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6000

(a)

ZZ” (ohm.cm )

5000

some samples show an inductive loop at low frequencies. As it has been widely reported [44,45], the high to middle frequency loops are attributed to the coating resistance and the charge transfer process. While the low frequencies inductive loop is due to the presence of metastable Mg+ ions during the anodic dissolution of magnesium alloy substrate. The metastable Mg+ ions converted to Mg2+ followed by the formation of gelatin compound of Mg(OH)2 on the surface [46]. The inductive loop is seen in D-CP, D-HA and P-CP coated samples in which the SBF solution may infiltrate the coating and reach the substrate, while the coatings with a better corrosion resistance did not show such an inductive loop (Fig. 5). In order to quantitatively analyze the EIS data, the experimental results (Fig. 5) were modeled by the electrical equivalent circuits shown in Figs. 6a and 6b. The equivalent circuit parameters derived from EIS measurements of uncoated and coated samples are presented in Table 3. In the equivalent circuits, Rs represents solution resistance. R1 is related to the coating resistance, R2 and RL are the charge transfer and inductive resistance, respectively. CPE impedance is equals to [Q (jω)n]−1 which is a constant phase element represents a non-ideal behavior of capacitive level. In terms of pure capacitor behavior n will be equal to 1, while in reality this number varies between 0 and 1 [47]. It was reported that by decreasing the n value, the surface of the specimen became rougher [47]. As it can be seen, the highest n belongs to bare AZ31 alloy; additionally pulse electrodeposited coatings have higher n compared to direct one. This may show that coatings obtained by pulse electrodeposition are neater than direct ones. The polarization resistance (Rpr) of EIS results was calculated by the following Eq. (5) [40]:

AZ31 D-CP D-HA D-HA-C0.25 D-HA-C0.5 D-HA-C1 D-HA-C1.5

4000

3000 2000 1000 0

-1000

0

3000

6000

9000

12000

15000

18000

ZZ’ (ohm.cm ) 6000

ZZ” (ohm.cm )

AZ31

(b)

5000

P-CP P-HA

P-HA-C0.25

4000

P-HA-C0.5

3000

P-HA-C1

P-HA-C1.5

2000

1000 0

-1000

0

3000

6000

9000

12000

15000

18000

ZZ’ (ohm.cm ) Fig. 5. Nyquist plot after 15 min immersion in SBF solution for uncoated AZ31 magnesuim alloy and different coatings conditions: (a) Direct Voltage, (b) Pulse Voltage.

(5)

Rpr = R1 + R2

As it can be seen, the pulse voltage deposited coatings have higher polarization resistance than that of direct ones. Moreover, HA-CNTs composite coatings always illustrate higher overall corrosion resistance, (R1 + R2), and lower CPE values in comparison to pure HA coatings which confirm their better anti-corrosive behavior. The following trend indicates the enhancing polarization resistance in both direct and pulse voltage deposited coatings which is in compliance with the other electrochemical measurements: AZ 31 < CP < HA < HA

C 0. 25 < HA

C 0. 5 < HA

C1. 5 < HA

C1

The overall corrosion resistance of P-HA-C1, as an optimized sample, is 16,530 Ω cm2 which increase the corrosion resistance of pulsed deposited HA sample and bare AZ31 alloy for 5381 and 15,522 Ω cm2, respectively. 3.2. Immersion testing Fig. 6. Equivalent circuits used to fit the EIS data: (a) with an inductive loop at low frequencies, (b) without Inductive loop at low frequencies.

This confirms again the homogeneity of the coatings obtained by the pulse deposition method. By adding 1% CNTs to the coatings (Figs. 4c and 4d), the number of pits declined. No apparent pit is observed on the surface of the P-HA-C1 specimen. Consequently, OCP and polarization tests results prove that the P-HA-C1 sample not only has the noble OCP in SBF solution compared to the other coatings conditions, but also it prevents the growth of stable corrosion pits in the anodic potentials during the polarization test. 3.1.3. Electrochemical Impedance Spectroscopy (EIS) measurements In order to further analyze the corrosion behavior of the samples, the electrochemical impedance spectroscopy tests were employed. The EIS spectra of AZ31 magnesium alloy and the coated samples, after 15 min immersion in SBF solution, are shown in Fig. 5. The EIS curves show two semi-circles at high and middle frequencies. Additionally,

3.2.1. pH determination According to some research results [48], the alkaline pH of the environment around the implant has a detrimental effect on living cells. Therefore, the pH change during immersion of samples in SBF is important to be determined. In this regard, the samples were soaked for 5 days in SBF solution at the temperature of 37 ± 1 °C. Fig. 7, indicates the variation of pH values of AZ31, D-HA, P-HA, D-HA-C1 and P-HA-C1 versus immersion time. On the one hand, AZ31 sample shows an immediate slight decrease in pH solution followed by a rapid increase during the immersion time. Once the bare specimen is immersed in SBF solution, it is corroded and the corrosion products are formed on the surface (Eqs. (1)–(4)) [49]. The OH- ions are consumed, while the Mg (OH)2 is formed on the surface of AZ31 leading to decrease in pH solution. The high amount of chloride in the SBF can convert the Mg(OH)2 to MgCl2 according to Eq. (6), in which MgCl2 is easily soluble in the body fluid as indicated in Eq. (7) [50]. Consequently, Cl- ions release which promotes further dissolution of Mg(OH)2 and corrosion of AZ31. This leads to dramatic increase in pH value in the further stages of immersion.

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Table 3 Extracted parameter from the EIS data after 15 min immersion of samples in SBF solution. Sample label AZ31 D-CP D-HA D-HA-C0.25 D-HA-C0.5 D-HA-C1 D-HA-C1.5 P-CP P-HA P-HA-C0.25 P-HA-C0.5 P-HA-C1 P-HA-C1.5

R1 (Ω cm2)

CPE1 (μF cm−2)

254 4137 5342 7134 8010 9721 9289 6236 8052 9114 10,109 12,521 12,243

9.4 6.34 4.85 5.12 3.69 2.47 3.56 5.314 3.63 4.52 2.91 1.62 3.89

P-HA-C1

8.4

D-HA-C1

n

1

R2 (Ω cm2)

CPE2 (μF cm−2)

754 1667 1851 2786 2835 2991 1833 2873 3097 2989 3046 4009 3455

299 258 243 198 217 182 231 234 277 273 197 131 158

0.96 0.71 0.78 0.69 0.77 0.79 0.70 0.84 0.90 0.85 0.82 0.88 0.81

P-HA

D-HA

MgCl2(s) →

PH

8 7.8 7.6

7.4

0

20

40 60 80 100 Immersion Time (hour)

120

Hydrogenn evolution amount (ml.cm-2)

Fig. 7. The pH values of SBF solution versus immersion time at 37 ± 1 °C for AZ31, pure hydroxyapatite and composite coatings with 1 wt% CNTs.

P-HA-C1

1

D-HA-C1

P-HA

D-HA

AZ31

0.8

0.6

0.4

0.2

0 0

20

40

60

80

100

2

0.98 0.50 0.51 0.55 0.53 0.57 0.56 0.59 0.60 0.52 0.53 0.59 0.58

RL (Ω cm2)

L1 (H/cm2)

Rpr (Ω cm2)

230.5 201.4 221.9 ___ ___ ___ ___ 217.1 ___ ___ ___ ___ ___

190.2 187.3 163.5 ___ ___ ___ ___ 159.8 ___ ___ ___ ___ ___

1008 5804 7193 9920 10,845 12,712 11,122 9109 11,149 12,103 13,155 16,530 15,698

− Mg(OH)2(s) + 2Cl− (aq) → MgCl2(s) + 2OH

AZ31

8.2

7.2

n

120

Immersion Time (hour) Fig. 8. Hydrogen evolution rate during immersion in SBF solution at 37 ± 1 °C for pure hydroxyapatite and composite coatings with 1 wt% CNTs.

Mg2+ (aq)

+

2Cl− (aq)

(6) (7)

On the other hand, the solution pH of the coated samples increase rapidly in the early stages of immersion. This is mainly due to the partial dissolution of pre-exist coating in SBF [51] and releasing calcium, phosphate and hydroxyl ions leading to pH increase. Moreover, during the immersion time, SBF solution might infiltrate into the interface between the coating and substrate through the defects resulting in corrosion of the substrate and an increase in pH value (Eq. (2)). Compared to the solution pH of AZ31 alloy, all the coated samples apparently control the pH increase showing that the coatings act as a barrier which slows down the accessibility of electrolyte to the substrate. Furthermore, the rate of pH increase is slower in pulse deposited and CNT added coatings compared to direct deposited and pure HA coatings, respectively. These indications confirm the results obtained from electrochemical experiments in previous parts. After about 60 h of immersion, the rate of increment in pH slows down for all the coated samples. According to the previous researches by increasing the pH, the hydroxyapatite tendency for dissolution in SBF is declined [42]. Moreover, new calcium phosphate coating may form on the sample which consumes the OH- ions. The formation of new deposits can slow down the degradation rate of the substrate by enhancing the coating resistance against corrosion. A slight decrease in solution pH of P-HA-C1 sample was observed after 100 h of immersion. This may show that the new coating formed on P-HA-C1 (consuming OH- ions), retards the dissolution of the existed coating as well as acting as a better barrier against the substrate corrosion (providing OH- ions) compared to the other samples. 3.2.2. Hydrogen evolution The results of the hydrogen evolution during immersion time are shown in Fig. 8. According to Eqs. (1) and (2), it can be seen that the corrosion of magnesium alloy follows by H2 gas generation. By comparing Fig. 7 and Fig. 8, it can be seen that the pH value and the hydrogen generation increased by soaking time with relatively the same trend for coated samples. The continuous release of H2 bubbles indicates that the cathodic and anodic reactions are active during the 5 days of immersion showing the corrosion of magnesium substrate in all the coated samples. The H2 generation is dramatically declined in the coated samples. In more detail, adding 1% CNTs to HA coating in pulse deposition method causes a tenfold decrease in the amount of hydrogen evolution compared to the bare AZ31 alloy. According to previous research [52], the reduction in hydrogen liberation could encourage the proliferation of bone cells and reduce the healing time, so P-HA-C1 sample has the most suitable condition as a degradable implant.

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Fig. 9. SEM images after 5 days immersion in SBF solution for pure hydroxyapatite and composite coatings with 1 wt% CNTs in different magnifications.

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Table 4 The EDS results of Fig. 9 at low magnification after 5 days immersion in SBF solution. Direct voltage Sample label

D-HA D-HA-C1

Pulse voltage

Ca/P ratio Before immersion

After immersion

1.48 1.53

1.58 1.66

Sample label

P-HA P-HA-C1

Ca/P ratio Before immersion

After immersion

1.52 1.55

1.60 1.67

Ca/P ratio than 1.67 stoichiometric HA. However, the new deposits formed after five days immersion in SBF solution have higher Ca/P ratios. It shows that the pre-exist coatings, decreased the corrosion of the magnesium substrate during the soaking time which is in agreement with electrochemical tests results. Moreover, the increase in Ca/P ratio in composite coatings are higher than pure hydroxyapatite coatings. It may concluded that more Ca2+ and PO43− ions were absorbed on the surface of composite coatings which were related to increasing the nucleation sites by functionalized CNTs. 4. Conclusion

3.2.3. Microstructural characterization The ability to enhance the apatite formation is considered as one of the major factors to define a bio-ceramic as a bioactive material. In this regard, the morphology of the samples after 5 days immersion in SBF solution at 37 ± 1 °C was studied and the results are illustrated in Fig. 9. It is worth mentioning that SBF solution used in this survey does not contain any biological matter such as bone forming cells. The coating structure of D-HA and P-HA samples after 5 days immersion are almost similar to that of before immersion (Fig. 5 of Part 1) at low magnification. However, the formation of new spherical deposits is seen on the main plate-like structure of these coatings at high magnification. According to previous reports [53,54], two phenomena occurred while the calcium phosphate coated samples were soaked in SBF solution: i) dissolution of pre-exist coating ii) nucleation and growth of new deposits by increasing the amount of ions in SBF near the coatings’ surface. It can be explained that the concentrations of calcium and phosphate ions in the SBF solution are relatively low, so that the structures of D-HA and P-HA samples remain intact, in low magnification. However, in CNTs added coatings, the new spherical deposits covered the pre-exist plate-like structure in both high and low magnification images in just 5 days immersion. This indicates that the increase of nucleation sites due to the functionalized carbon nano-tubes can compensate the low concentration of calcium and phosphate ions leading to a decrease in nucleation formation energy of new deposits. It can also be added that the new layer formed on P-HA-C1 sample has finer structure compared to D-HA-C1 one which can be attributed to their coatings’ morphology before immersion (Fig. 6 of Part 1). The P-HA-C1 sample has a finer structure with a higher specific area that can provide more sites for absorption of Ca2+, PO43− and OH- ions from SBF solution. The formation of new deposits during 5 days immersion, proved our claim in pH determination section. Xin and his coworkers [55] showed a different trend. They claimed that in the early stages of immersion (4–7 days), the new crystals formed and grew with preferential orientation due to the low concertation of Ca and P ions in this stage. In the later stage (more than 7 days), the concentration of Ca and P ions were increased which provided more nucleation sites on the coating surface followed by spherical crystal formation and growth instead of deposits with preferable orientation. However, in our research the new spherical deposits are formed and covered the surface in HA-CNTs composite coatings, in just 5 days. This shows that HA-CNTs composite coating is bioactive and could induce the surface bio-mineralization in SBF solution.

Based on the experimental results, it was concluded that pure HA and HA-CNTs composite coating on AZ31 magnesium alloy could reduce its corrosion in SBF solution. The following observation were also made. 1. The pulsed-voltage electrodeposited coatings show better corrosion resistance tested by OCP, Tafel polarization and EIS methods compared to the direct electrodeposited coatings, because of their fine, dense and uniform structure. 2. Adding functionalized CNTs to HA structure improves the corrosion resistance of the coating by reducing the corrosive solution access to the substrate. This happens as a result of the increase in HA nucleation sits due to carboxyl functional groups of CNTs, higher crystallinity and bridging mechanism of CNTs between HA structure leading to densify the composite coating. 3. Pulsed electrodeposited HA having 1% CNTs, was the optimized condition. To assess it quantitatively, Ecorr of the AZ31 Mg alloy increased from −1.57 V vs. Ag/AgCl to –1.40 and −1.30 V vs. Ag/ AgCl for P-HA and P-HA-C1 coatings, respectively. In addition, corrosion current density of P-HA-C1 coated sample (0.72 µA/cm2) is respectively one order of magnitude (2.51 µA/cm2) and two orders of magnitude (44.2 µA/cm2) lower than P-HA coated sample and bare one. 4. Pulsed electrodeposited HA having 1% CNTs caused a tenfold decrease in the amount of hydrogen generation of bare AZ31 magnesium alloy and significantly reduced the pH change in the vicinity of the sample, soak in SBF solution. 5. The addition of CNTs did not affect the apatite-forming ability on the surface of HA-CNTs composite coating. Moreover, the present results reveal that these coatings are bone bioactive and hence favor osseointegration. References

3.2.4. Compositional analysis Energy Dispersive Spectroscopy (EDS) test was performed on the low magnification images of Fig. 9 which indicates the Ca/P ratio after 5 days of immersion. These results as well as the Ca/P ratio determined before the immersion in part 1 of this study, are presented in Table 4. It can be seen that the Ca/P ratio of the surface of the coatings, increased after 5 days of immersion in SBF solution. In Part 1 of this survey, it was claimed that during the electrodeposition process, Mg ions might release due to the high activity and corrosion rate of Mg substrate. These ions can substitute for Ca ions in the coatings which result in the lower 11193

[1] G. Ciobanu, M. Harja, Cerium-doped hydroxyapatite/collagen coatings on titanium for bone implants, Ceram. Int. (2018). [2] Y.-C. Liu, G.S. Lin, J.-Y. Wang, C.-S. Cheng, Y.-C. Yang, B.-S. Lee, K.-L. Tung, Synthesis and characterization of porous hydroxyapatite coatings deposited on titanium by flame spraying, Surf. Coat. Technol. (2018). [3] J. Yang, K. Zhang, K. Que, S. Hou, Z. Chen, Y. Li, Y. Wang, Y. Song, B. Guan, W. Zhang, Surface modification of titanium with hydroxyapatite layer induced by phase-transited lysozyme coating, Mater. Sci. Eng. C 92 (2018) 206–215. [4] X. Zhang, G. Xiao, B. Liu, C. Jiang, N. Li, Y. Lu, The formation of hydroxyapatite layer onto hopeite coating on stainless steel substrate, Corros. Sci. 111 (2016) 216–229. [5] A. Bekmurzayeva, W.J. Duncanson, H.S. Azevedo, D. Kanayeva, Surface modification of stainless steel for biomedical applications: revisiting a century-old material, Mater. Sci. Eng. C (2018). [6] S. Heise, T. Wirth, M. Höhlinger, Y.T. Hernández, J.A.R. Ortiz, V. Wagener, S. Virtanen, A.R. Boccaccini, Electrophoretic deposition of chitosan/bioactive glass/ silica coatings on stainless steel and WE43 Mg alloy substrates, Surf. Coat. Technol. 344 (2018) 553–563. [7] M. Mardali, H.R. SalimiJazi, F. Karimzadeh, B. Luthringer, C. Blawert, S. Labbaf, Comparative study on microstructure and corrosion behavior of nanostructured hydroxyapatite coatings deposited by high velocity oxygen fuel and flame spraying on AZ61 magnesium based substrates, Appl. Surf. Sci. (2018). [8] W. Yu, R. Sun, Z. Guo, Z. Wang, Y. He, G. Lu, P. Chen, K. Chen, Novel fluoridated hydroxyapatite/MAO composite coating on AZ31B magnesium alloy for biomedical application, Appl. Surf. Sci. (2018).

Ceramics International 45 (2019) 11186–11194

D. Khazeni, et al. [9] S. Tabasum, A. Noreen, A. Kanwal, M. Zuber, M.N. Anjum, K.M. Zia, International journal of biological macromolecules glycoproteins functionalized natural and synthetic polymers for prospective biomedical applications: a review, Int. J. Biol. Macromol. 98 (2017) 748–776, https://doi.org/10.1016/j.ijbiomac.2017.01.078. [10] J.P. Santerre, K. Woodhouse, G. Laroche, R.S. Labow, Understanding the biodegradation of polyurethanes: from classical implants to tissue engineering materials, Biomaterials 26 (2005) 7457–7470, https://doi.org/10.1016/j.biomaterials.2005. 05.079. [11] J.V. Rau, I. Antoniac, M. Filipescu, C. Cotrut, M. Fosca, L.C. Nistor, R. Birjega, M. Dinescu, Hydroxyapatite coatings on Mg-Ca alloy prepared by pulsed laser deposition: properties and corrosion resistance in simulated body fluid, Ceram. Int. (2018). [12] W. Huang, B. Xu, W. Yang, K. Zhang, Y. Chen, X. Yin, Y. Liu, Z. Ni, F. Pei, Corrosion behavior and biocompatibility of hydroxyapatite/magnesium phosphate/zinc phosphate composite coating deposited on AZ31 alloy, Surf. Coat. Technol. 326 (2017) 270–280. [13] C. Prakash, S. Singh, B.S. Pabla, M.S. Uddin, Synthesis, characterization, corrosion and bioactivity investigation of nano-HA coating deposited on biodegradable MgZn-Mn alloy, Surf. Coat. Technol. 346 (2018) 9–18, https://doi.org/10.1016/j. surfcoat.2018.04.035. [14] G. Song, Recent progress in corrosion and protection of magnesium alloys, Adv. Eng. Mater. 7 (2005) 563–586. [15] S. Agarwal, J. Curtin, B. Duffy, S. Jaiswal, Biodegradable magnesium alloys for orthopaedic applications: a review on corrosion, biocompatibility and surface modifications, Mater. Sci. Eng. C 68 (2016) 948–963. [16] N.T. Kirkland, N. Birbilis, M.P. Staiger, Acta Biomaterialia Assessing the corrosion of biodegradable magnesium implants: a critical review of current methodologies and their limitations, Acta Biomater. 8 (2012) 925–936, https://doi.org/10.1016/j. actbio.2011.11.014. [17] M. Esmaily, J.E. Svensson, S. Fajardo, N. Birbilis, G.S. Frankel, S. Virtanen, R. Arrabal, S. Thomas, L.G. Johansson, Fundamentals and advances in magnesium alloy corrosion, Prog. Mater. Sci. 89 (2017) 92–193. [18] M.A. Leeflang, J.S. Dzwonczyk, J. Zhou, J. Duszczyk, Long-term biodegradation and associated hydrogen evolution of duplex-structured Mg–Li–Al–(RE) alloys and their mechanical properties, Mater. Sci. Eng. B 176 (2011) 1741–1745. [19] G.S. Frankel, A. Samaniego, N. Birbilis, Evolution of hydrogen at dissolving magnesium surfaces, Corros. Sci. 70 (2013) 104–111. [20] H. Hornberger, S. Virtanen, A.R. Boccaccini, Biomedical coatings on magnesium alloys - a review, Acta Biomater. 8 (2012) 2442–2455, https://doi.org/10.1016/j. actbio.2012.04.012. [21] H. Tang, Y. Han, T. Wu, W. Tao, X. Jian, Y. Wu, F. Xu, Synthesis and properties of hydroxyapatite-containing coating on AZ31 magnesium alloy by micro-arc oxidation, Appl. Surf. Sci. 400 (2017) 391–404, https://doi.org/10.1016/j.apsusc.2016. 12.216. [22] H. Pan, H. Yang, X. Tang, J. Niu, Z. Xiang, Y. Song, W. Lu, Effect of Ca/P ratio on the structural and corrosion properties of biomimetic Ca[sbnd]P coatings on ZK60 magnesium alloy, Mater. Sci. Eng. C 72 (2017) 676–681, https://doi.org/10.1016/j. msec.2016.11.132. [23] S. Shen, S. Cai, Y. Li, R. Ling, F. Zhang, G. Xu, F. Wang, Microwave aqueous synthesis of hydroxyapatite bilayer coating on magnesium alloy for orthopedic application, Chem. Eng. J. 309 (2017) 278–287. [24] S. Shen, S. Cai, X. Bao, P. Xu, Y. Li, S. Jiang, G. Xu, Biomimetic fluoridated hydroxyapatite coating with micron/nano-topography on magnesium alloy for orthopaedic application, Chem. Eng. J. 339 (2018) 7–13, https://doi.org/10.1016/j. cej.2018.01.083. [25] A. Szcześ, L. Hołysz, E. Chibowski, Synthesis of hydroxyapatite for biomedical applications, Adv. Colloid Interface Sci. 249 (2017) 321–330. [26] H. Li, X. Song, B. Li, J. Kang, C. Liang, H. Wang, Z. Yu, Z. Qiao, Carbon nanotubereinforced mesoporous hydroxyapatite composites with excellent mechanical and biological properties for bone replacement material application, Mater. Sci. Eng. C 77 (2017) 1078–1087, https://doi.org/10.1016/j.msec.2017.04.048. [27] M. Ding, N. Sahebgharani, F. Musharavati, F. Jaber, E. Zalnezhad, G.H. Yoon, Synthesis and properties of HA/ZnO/CNT nanocomposite, Ceram. Int. 44 (2018) 7746–7753. [28] V.E. Kagan, N.V. Konduru, W. Feng, B.L. Allen, J. Conroy, Y. Volkov, I.I. Vlasova, N.A. Belikova, N. Yanamala, A. Kapralov, Carbon nanotubes degraded by neutrophil myeloperoxidase induce less pulmonary inflammation, Nat. Nanotechnol. 5 (2010) 354. [29] H.-X. Ren, X. Chen, J.-H. Liu, N. Gu, X.-J. Huang, Toxicity of single-walled carbon nanotube: how we were wrong? Mater. Today 13 (2010) 6–8. [30] M.A. Hussain, M.A. Kabir, A.K. Sood, On the cytotoxicity of carbon nanotubes, Curr. Sci. 96 (2009). [31] Y. Usui, K. Aoki, N. Narita, N. Murakami, I. Nakamura, K. Nakamura, N. Ishigaki, H. Yamazaki, H. Horiuchi, H. Kato, Carbon nanotubes with high bone‐tissue compatibility and bone‐formation acceleration effects, Small 4 (2008) 240–246.

[32] S. Giannona, I. Firkowska, J. Rojas-Chapana, M. Giersig, Vertically aligned carbon nanotubes as cytocompatible material for enhanced adhesion and proliferation of osteoblast-like cells, J. Nanosci. Nanotechnol. 7 (2007) 1679–1683. [33] A.O. Lobo, M.A.F. Corat, E.F. Antunes, M.B.S. Palma, C. Pacheco-Soares, E.E. Garcia, E.J. Corat, An evaluation of cell proliferation and adhesion on vertically-aligned multi-walled carbon nanotube films, Carbon N. Y. 48 (2010) 245–254. [34] D. Gopi, E. Shinyjoy, M. Sekar, M. Surendiran, L. Kavitha, T.S. Sampath Kumar, Development of carbon nanotubes reinforced hydroxyapatite composite coatings on titanium by electrodeposition method, Corros. Sci. 73 (2013) 321–330, https://doi. org/10.1016/j.corsci.2013.04.021. [35] X. Pei, Y. Zeng, R. He, Z. Li, L. Tian, J. Wang, Q. Wan, X. Li, H. Bao, Single-walled carbon nanotubes/hydroxyapatite coatings on titanium obtained by electrochemical deposition, Appl. Surf. Sci. 295 (2014) 71–80. [36] R. Singh, D. Pantarotto, L. Lacerda, G. Pastorin, C. Klumpp, M. Prato, A. Bianco, K. Kostarelos, Tissue biodistribution and blood clearance rates of intravenously administered carbon nanotube radiotracers, Proc. Natl. Acad. Sci. USA 103 (2006) 3357–3362. [37] K. Kostarelos, Carbon nanotubes: fibrillar pharmacology, Nat. Mater. 9 (2010) 793. [38] S. Shadanbaz, G.J. Dias, Calcium phosphate coatings on magnesium alloys for biomedical applications: a review, Acta Biomater. 8 (2012) 20–30, https://doi.org/ 10.1016/j.actbio.2011.10.016. [39] E.C. Meng, S.K. Guan, H.X. Wang, L.G. Wang, S.J. Zhu, J.H. Hu, C.X. Ren, J.H. Gao, Y.S. Feng, Effect of electrodeposition modes on surface characteristics and corrosion properties of fluorine-doped hydroxyapatite coatings on Mg-Zn-Ca alloy, Appl. Surf. Sci. 257 (2011) 4811–4816, https://doi.org/10.1016/j.apsusc.2010.12.073. [40] P. Wu, Z. Zhang, F. Xu, K. Deng, K. Nie, R. Gao, Effect of duty cycle on preparation and corrosion behavior of electrodeposited calcium phosphate coatings on AZ91, Appl. Surf. Sci. 426 (2017) 418–426. [41] Y. Shangguan, P. Wan, L. Tan, X. Fan, L. Qin, K. Yang, Investigation of the inner corrosion layer formed in pulse electrodeposition coating on Mg-Sr alloy and corresponding degradation behavior, J. Colloid Interface Sci. 481 (2016) 1–12, https:// doi.org/10.1016/j.jcis.2016.07.032. [42] M. Canillas, P. Pena, A.H. De Aza, M.A. Rodríguez, Calcium phosphates for biomedical applications, Bol. La Soc. Esp. Ceram. Vidr. 56 (2017) 91–112, https://doi. org/10.1016/j.bsecv.2017.05.001. [43] L. Sun, C.C. Berndt, K.A. Gross, A. Kucuk, Material fundamentals and clinical performance of plasma‐sprayed hydroxyapatite coatings: a review, J. Biomed. Mater. Res. Off. J. Soc. Biomater. Jpn. Soc. Biomater. Aust. Soc. Biomater. Korean Soc. Biomater. 58 (2001) 570–592. [44] T.S. Lim, H.S. Ryu, S.-H. Hong, Electrochemical corrosion properties of CeO2containing coatings on AZ31 magnesium alloys prepared by plasma electrolytic oxidation, Corros. Sci. 62 (2012) 104–111. [45] L.-Y. Cui, S.-D. Gao, P.-P. Li, R.-C. Zeng, F. Zhang, S.-Q. Li, E.-H. Han, Corrosion resistance of a self-healing micro-arc oxidation/polymethyltrimethoxysilane composite coating on magnesium alloy AZ31, Corros. Sci. 118 (2017) 84–95. [46] Y. Song, D. Shan, R. Chen, F. Zhang, E.-H. Han, Biodegradable behaviors of AZ31 magnesium alloy in simulated body fluid, Mater. Sci. Eng. C 29 (2009) 1039–1045. [47] Y. Su, L. Niu, Y. Lu, J. Lian, G. Li, Preparation and corrosion behavior of calcium phosphate and hydroxyapatite conversion coatings on AM60 magnesium alloy, J. Electrochem. Soc. 160 (2013) C536–C541. [48] J. Hu, F. Wu, S. Feng, J. Xu, Z. Xu, Y. Chen, T. Tang, X. Weng, X. Zhou, A convenient ratiomeric pH probe and its application for monitoring pH change in living cells, Sens. Actuators B Chem. 196 (2014) 194–202. [49] J. Liao, M. Hotta, Y. Mori, Improved corrosion resistance of a high-strength Mg-AlMn-Ca magnesium alloy made by rapid solidification powder metallurgy, Mater. Sci. Eng. A 544 (2012) 10–20, https://doi.org/10.1016/j.msea.2012.02.046. [50] H. Altun, S. Sen, Studies on the influence of chloride ion concentration and pH on the corrosion and electrochemical behaviour of AZ63 magnesium alloy, Mater. Des. 25 (2004) 637–643. [51] D.T.M. Thanh, P.T. Nam, N.T. Phuong, L.X. Que, N. Van Anh, T. Hoang, T. Dai Lam, Controlling the electrodeposition, morphology and structure of hydroxyapatite coating on 316L stainless steel, Mater. Sci. Eng. C 33 (2013) 2037–2045. [52] R. Rojaee, M. Fathi, K. Raeissi, A. Sharifnabi, Biodegradation assessment of nanostructured fluoridated hydroxyapatite coatings on biomedical grade magnesium alloy, Ceram. Int. 40 (2014) 15149–15158. [53] Q. Zhang, J. Chen, J. Feng, Y. Cao, C. Deng, X. Zhang, Dissolution and mineralization behaviors of HA coatings, Biomaterials 24 (2003) 4741–4748, https://doi. org/10.1016/S0142-9612(03)00371-5. [54] Y. Cai, S. Zhang, X. Zeng, Y. Wang, M. Qian, W. Weng, Improvement of bioactivity with magnesium and fluorine ions incorporated hydroxyapatite coatings via sol–gel deposition on Ti6Al4V alloys, Thin Solid Films 517 (2009) 5347–5351. [55] F. Xin, C. Jian, J.-P. Zou, W. Qian, Z.-C. Zhou, J.-M. Ruan, Bone-like apatite formation on HA/316L stainless steel composite surface in simulated body fluid, Trans. Nonferr. Met. Soc. China 19 (2009) 347–352.

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