J Mater Sci: Mater Med DOI 10.1007/s10856-013-5127-7

Nano-hydroxyapatite reinforced AZ31 magnesium alloy by friction stir processing: a solid state processing for biodegradable metal matrix composites B. Ratna Sunil • T. S. Sampath Kumar Uday Chakkingal • V. Nandakumar • Mukesh Doble

•

Received: 29 May 2013 / Accepted: 13 December 2013 Ó Springer Science+Business Media New York 2013

Abstract Friction stir processing (FSP) was successfully adopted to fabricate nano-hydroxyapatite (nHA) reinforced AZ31 magnesium alloy composite as well as to achieve fine grain structure. The combined effect of grain refinement and the presence of embedded nHA particles on enhancing the biomineralization and controlling the degradation of magnesium were studied. Grain refinement from 56 to *4 and 2 lm was observed at the stir zones of FSP AZ31 and AZ31–nHA composite respectively. The immersion studies in super saturated simulated body fluid (SBF 59) for 24 h suggest that the increased wettability due to fine grain structure and nHA particles present in the AZ31–nHA composite initiated heterogeneous nucleation which favored the early nucleation and growth of calciumphosphate mineral phase. The nHA particles as nucleation sites initiated rapid biomineralization in the composite. After 72 h of immersion the degradation due to localized pitting was observed to be reduced by enhanced biomineralization in both the FSPed AZ31 and the composite. Also, best corrosion behavior was observed for the composite before and after immersion test. MTT assay using rat skeletal muscle (L6) cells showed negligible toxicity for all the processed and unprocessed samples. However, cell adhesion was observed to be more on the composite due to the small grain size and incorporated nHA.

B. Ratna Sunil  T. S. Sampath Kumar (&)  U. Chakkingal Department of Metallurgical and Materials Engineering, Indian Institute of Technology Madras, Chennai 600036, India e-mail: [email protected] V. Nandakumar  M. Doble Department of Biotechnology, Indian Institute of Technology Madras, Chennai 600036, India

1 Introduction The study of magnesium and its alloys as degradable implants is one of the promising research topics in the area of biomaterials. Prime interest behind tailoring the magnesium for biomaterial application is their load bearing capacity and good mechanical properties near to the natural human bone. The density, elastic modulus and yield strength of magnesium are closer to the bone tissue compared to other metallic implants, which can minimize or avoid the stress shielding effect [1, 2]. But magnesium undergoes rapid corrosion in physiological conditions and produces magnesium hydroxide (Mg(OH)2) leading to the evolution of hydrogen gas. In presence of chloride ions, magnesium hydroxide starts to convert into magnesium chloride, which dissolves quickly in biological solution resulting in localized severe pitting. This leads to an increase in the dissolution rate of magnesium compared to the rate of new bone formation [3–5]. Although, the corrosion products are nontoxic, mechanical properties and rate of tissue healing are severely affected due to uncontrolled rapid degradation of magnesium. Also, the rate of hydrogen evolution affects the healing process due to lower cell activities at the material tissue interface [6]. Therefore, the need for reducing the degradation rate is essential in developing biodegradable implants based on magnesium biocorrosion. Preparing new alloys and composites, developing surface coatings and modifying the microstructure are the few methods employed to alter the degradation rate of magnesium based materials [2, 6–10]. AZ series (Mg–Al–Zn) of alloys are common magnesium alloys with superior mechanical properties, corrosion resistance and good castability [11]. With low Al content, good mechanical properties and corrosion resistance, AZ31 magnesium alloy can be a suitable biodegradable material.

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Fig. 1 Friction stir processing: a schematic representation, b photographs of the sample before processing and c during the process

On the other hand, hydroxyapatite (HA), a calcium phosphate mineral that resembles the natural bone mineral phase has emerged as a promising bioceramic material for the past two decades due to its excellent biocompatibility, bioactivity and bone bonding ability [12]. Coating the surface of the implants with HA to improve bioactivity and osseointegration is a well known method. However, for degradable metals like magnesium, the quality of the coating determines the degradation behavior of the substrate. If the subsurface is exposed to the corroding environment due to defects in the coating, the rate of degradation will be uncontrolled at the exposed locations. The bonding between the coating and the substrate will also be affected. Since, the degradation of magnesium initiates from the surface, incorporating HA into the surface and sub-surface layers rather can be a promising technique to control the rate of degradation than providing surface coatings. In the present study, the grain refinement and dispersion of nano-hydroxyapatite (nHA) particles in AZ31 magnesium alloy was achieved by adopting friction stir processing (FSP), a new solid-state processing technique developed on the basic principles of friction stir welding [13]. The principle behind the grain refinement in FSP has been explained by Mishra et al. [13, 14]. During FSP, the

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cylindrical rotating tool consisting of a small pin as shown in the Fig. 1a is inserted into the material surface and plunged along desired length causes dynamic recrystallization due to intense plastic deformation within the stir zone [13] and leads to grain refinement. The stirring action of the FSP tool can be used to incorporate and distribute particles of secondary phase in order to develop fine grained metal matrix composites (MMC). For the first time, Mishra et al. [15] successfully fabricated 5083Albased SiC reinforced surface composite using FSP. Significant work has been done in developing MMC using oxides or carbides as the dispersing media and aluminum or copper as the matrix. Most of the reported work has been on studying the microstructural evolution, hardness, corrosion and wear properties [16–22]. However, reports concerning the fabrication of magnesium based composites by FSP are limited. In particular, fabricating the composites has not been investigated for biomedical applications. In the present paper, we demonstrate the fabrication of AZ31–nHA composites using FSP for degradable hard tissue applications. The role of microstructure and incorporated nHA particles on biomineralization and degradation has been studied. Cytotoxicity of the samples using rat skeletal muscle (L6) cells was also investigated.

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2 Experimental procedure 2.1 Materials and processing Commercially available AZ31 magnesium alloy (Exclusive Magnesium, Hyderabad, India) sheets (2.9 %Al, 0.88 %Zn, 0.001 %Fe, 0.02 %Mn and remaining being Mg) of size 150 9 100 9 4 mm were annealed at 613 K. The FSP tool is made of hardened H-13 tool steel with a shoulder diameter of 15 mm and a tapered pin with a diameter varying from 5 to 3 mm over 2.7 mm length. Processing parameters were optimized to achieve defect free processed samples. The rotating tool was traversed at a speed of 6 mm/min with a rotating speed of 1,200 rpm along the traverse axis of the work piece by applying 5,000 N load and the processed sample was coded as FSP AZ31. The grain size analysis before and after FSP was done by linear intercept method in which a set of lines are drawn randomly on the micrograph of known magnification and the average grain size is obtained by dividing the total length of the line by number of grains intercepted by the line [23]. Same processing parameters were adopted to produce the composite. A shallow groove of 1 mm width and 2 mm depth was produced on the surface of annealed AZ31 (AZ31) and the groove was filled with nHA powder. Lab prepared nHA powder was used in the present study. The synthesizing process of nHA by microwave irradiation was as reported by Rameshbabu et al. [24]. The next step was plunging the rotating tool by the pin into the sheet for stirring and producing the composite (named as AZ31–nHA). Figure 1 shows the schematic representation of the process and the photographs of the samples before and during FSP. 2.2 Characterization The samples were mechanically polished and etched with picric acid reagent (5 ml acetic acid, 5 gm picric acid, 10 ml water and 100 ml ethyl alcohol for 20–60 s). Microstructural observations were carried out using optical microscopy (Vertimet-CP, Chennai Metco, India), scanning electron microscopy (SEM, FEI Quanta 200, Netherland) operated at 30 kV and transmission electron microscopy (TEM, Philips CM12, Holland) operated at 120 kV. For TEM observations, thin samples were cut from the stir zone and were mechanically thinned to a thickness of 100 lm using different graded emery papers, and then electro-polished using a twin-jet electro-polishing facility using a solution mixture of 1 % perchloric acid and 99 % ethanol. 2.3 Dissolution of iron into the work piece About 25 mg of samples cut from the stir zone and unprocessed regions were dissolved in 2 % HNO3 aqueous

solution and the solution composition was evaluated by inductively coupled plasma atomic emission spectroscopy (ICP-AES, Optima 5300DV, Perkin Elmer, USA) method to study the possible dissolution of iron from the FSP tool into the work piece. 2.4 Wettability Wettability of the samples was investigated by measuring the contact angles. Before measuring the contact angles (Easy DROP, KRUSS, Germany), all the samples were mechanically polished using emery papers up to 2,000 grade. Measurements were obtained on the sample surfaces using distilled water as the solvent at five different locations under ambient conditions. The surface energy (Es) has been calculated from the contact angles using the following equation [25]. Es ¼ Evl cos h

ð1Þ

where Evl is the surface energy between water and air under ambient condition, (i.e., 72.8 mJ/m2 at 20 °C) for pure water and h is the static contact angle. 2.5 In vitro bioactivity in supersaturated SBF Bioactivity was investigated by immersing the specimens (of size 10 9 10 9 5 mm cut from the stir zone) in super saturated simulated body fluid (SBF 59) kept in a constant water bath at a temperature of 37 °C for 72 h. Ionic concentration of SBF is similar to those of human extracellular fluid and the preparation methodology was as reported by Kokubo and Takadama [26]. The ion concentration of SBF 59 used in the present study is listed in Table 1. Each sample was immersed in 50 ml of SBF solution (the ratio of the SBF volume to the sample apparent surface area is more than 1:10).The samples were removed from the SBF 59 and rinsed with de-ionized water, and subjected to different characterizations. The specimens after immersion Table 1 Ion concentrations of SBF 59 used in the present study Ion

Ion concentrations (mM) Blood plasma

SBF

SBF 59

Na?

142

142

710

K?

5

5

25

Mg2?

1.5

1.5

7.5

Ca2?

2.5

2.5

12.5

-

Cl

103

147.8

739

HCO3-

27

4.2

21

HPO42-

1.0

1.0

5.0

SO42-

0.5

0.5

2.5

pH

7.2–7.4

7.4

7.4

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test were characterized by X-ray powder diffractometer (D8 DISCOVER, Bruker, USA) with Cu Ka radiation ˚ ) at a scanning rate of 1 step/s and step size of (k = 1.54 A 0.1 °/step. The surface morphology was studied using SEM and the elemental composition of the surfaces was investigated by energy dispersive X-ray (EDS) analysis.

2.6.1 Degradation in immersion test The samples were immersed in SBF 59 (each sample in a volume of 500 ml) at 37 °C for 1, 2 and 3 days to measure the weight loss (weight loss = (weight before immersion - weight after cleaning the corrosion products)/surface area) [27]. The surface corrosion products were removed from the samples in boiling solution of chromic acid (180 g/1 l of de-ionized water) and then rinsed with ethanol and dried in air before measuring the weight loss. pH variation in the solution for 1, 2 and 3 days of immersion was measured (pH 700, Eutech instruments, Singapore) to observe the change of ionic concentration during the degradation of the samples. The corrosion rate (CR, mm/year) of the samples based on weight loss was calculated using the following equation [28]: ð2Þ

where W is the weight loss (g), A is the original surface area of the sample (cm2), T is the immersion time (h) and D is the density of the sample (g/cm3). In the present study we considered that the density is 1.74 g/cm3 for all the samples. 2.6.2 Electrochemical test Potentiodynamic polarization (Model K0235, EG&G Princeton Applied Research, USA) tests were conducted in SBF 59 solution to determine the corrosion rates of the samples. Experiments were conducted on the samples before and after biomineralization study. The counter electrode was made of graphite and the reference electrode was a saturated calomel electrode (SCE). One square cm area of the working electrode (sample) was exposed to the solution. All tests were conducted at room temperature (30 °C). Prior to the beginning of the polarization tests, the samples (of before immersion test) were metallographically polished using emery sheets up to 2,000 grade, ultrasonically cleaned with ethanol and the experiments were conducted immediately. The samples were fixed in the electrochemical cell using an O-ring and kept in the solution for 30 min to establish the open circuit potential (OCP). Immediately after OCP, without any further delay, voltage and corrosion current density measurements were

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CR ðmils=yearÞ ¼ 0:129  a  Icorr = n D

ð3Þ

where CR is the corrosion rate, a is the molar mass (for magnesium 24.3 g/mol), Icorr is the corrosion current density (lA/cm2), n is the valance and D is the density (1.74 gm/cm3). The obtained CR was converted into mm/ year by considering 1 mils/year equal to 0.0254 mm/year.

2.6 Corrosion behavior

CR ¼ 8:76  104 W = ATD

done between the potentials -2.2 and -0.2 V with a scanning rate of 5 mV/s on the electrochemical station. Afterwards, the corrosion rate was calculated using the following equation [29].

2.7 Cytotoxicity and cell adhesion The cytotoxicity of the samples (of size 10 9 10 9 1 mm cut from the central part of the stir zone of the samples) after exposure to rat skeletal muscle (L6) cells (National Centre for Cell Sciences, Pune, India) for 72 h was quantitatively determined using MTT [3-(4,5-181dimethylthiazole-2-yl)-2,5-diphenyl tetrazolium bromide] assay [30]. Different samples were placed into the wells of a 96-well culture plate and each well was seeded with 100 ll of cell suspension containing 1 9 105 cells. The cell seeded samples were incubated at 37 °C for 24, 48 and 72 h in a humidified atmosphere containing 5 % of CO2. 150 ll of culture medium was added to each well. 20 ll (5 mg/ml in PBS) of MTT solution was added to each well followed by incubation at 37 °C for 4 h. Dimethyl sulfoxide (DMSO) was added to the incubated samples and the absorbance value was measured at 570 nm with reference at 650 nm using a microplate reader (Bio-Rad, Richmond, CA, USA), with non-seeded wells set as blank. The experiments were carried in triplicate. Cell viability was calculated as the percentage relative to the control (standard polystyrene tissue culture plates) using the following equation. % Cell viability ¼ ðmean optical density =control optical densityÞ  100

ð4Þ

For cell adhesion study, the samples were sterilized with 70 % of ethanol and placed into the wells of a 96 well plate. About 1 9 105 cells were seeded on the surface of the samples and 200 ll of the culture media was added along the sides of the wall and incubated at 37 °C in humidified atmosphere containing 5 % of CO2 for 24 h. The culture media was removed and the samples were rinsed with PBS. Then the cells were fixed on the samples by adding 2.5 % of glutaraldehyde followed by rinsing with PBS. They were then dehydrated using various alcohol-water mixtures and immersed in iso-amylacetate for 2 min and dried. Samples with cells were then sputter coated with gold and observed using SEM to study the attachment and proliferation of the cells.

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2.8 Statistical analysis Statistical analysis was conducted using one-way ANOVA analysis to evaluate the difference in cell viability of the samples. A value of P \ 0.05 was considered to be statistically significant. The data was shown as the mean ± standard deviation.

3 Results and discussion 3.1 Microstructure Figure 2 shows the optical microscope images at the cross sections of the samples and TEM micrograph of nHA powder used in the present study. Melting of AZ31 magnesium alloy does not occur during FSP (i.e. temperature reached is lower than 650 °C, which is the melting temperature of magnesium) [14, 20]. Nano-HA is stable at this temperature as reported by Rameshbabu et al. [24]. The average grain size of annealed AZ31 was measured as 56 lm. The grain refinement was achieved up to 4 lm in FSP AZ31 and 2 lm in AZ31–nHA composite as shown in SEM observations (Fig. 3a, b). It can be found from these images that FSP leads to significant grain refinement. The

dispersed nHA particles appeared in the stir zone and are marked with white arrows (Fig. 3b). The corresponding EDS analysis clearly indicates the presence of calcium and phosphorous suggesting the incorporation of nHA into AZ31. The presence of agglomerated nHA was found at many locations (Fig. 3c) in AZ31–nHA composite. Even though, the width and depth of the groove on the work piece filled with nHA powder before FSP is 1 mm and 2 mm respectively, the nHA powder is not limited to 1 mm width after FSP, but is actually dispersed and distributed throughout the stir zone due to the plastic flow of the material within the stir zone that is equal to a width of FSP tool shoulder diameter (15 mm in the present study) and a depth of the groove (2 mm in the present study). Also, the length of the stir zone can be designed as per requirements (70 mm in the present study). The TEM images (Fig. 3d, e) shows the typical small grains in FSPed samples and embedded nHA individual crystals in the fine grains of AZ31–nHA composite (Fig. 3f). 3.2 Dissolution of iron into the work piece Presence of iron in magnesium influences the corrosion rate. If the Fe dissolution is more than about 0.015 % [7], the corrosion rate is abnormally increased to higher levels.

Fig. 2 Microstructural observatios before and after processing: a optical image of AZ31, b TEM image of nano-hydroxyapatite powder and corresponding SAED pattern, c optical image of FSP AZ31 and d AZ31–nHA

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Fig. 3 Microstructural observations: SEM images at the cross section of a FSP AZ31, b AZ31–nHA, c distribution of nHA particles in low magnification and TEM images of d FSP AZ31, e AZ31–nHA and

d individual HA crystals in AZ31–nHA composite represented with white arrows

Table 2 Elemental composition of the samples after and before FSP obtained from ICP-AES

angles for both the samples indicate that the samples are hydrophilic. Figure 4 shows typical photographs of water droplets on the surface of the samples and the calculated surface energies. The surface energies of FSP AZ31 and AZ31–nHA samples have been found to be higher compared to AZ31. Surface energy, which characterizes the wettability of a material, plays an important role in material–tissue interactions. Therefore, hydrophilic surfaces are favorable compared to hydrophobic for the adsorption of specific proteins which help to improve the cell activities at the implant surface [31]. Wettability of the materials can be manipulated directly by varying the surface properties especially by altering the surface energy by inducing grain refinement and increasing the surface roughness. In the present study, the influence of the surface roughness was negligible as all the samples were mechanically polished to the same level of roughness using emery papers before performing the measurements. Therefore, the variation in contact angles is attributed to the difference in grain sizes only. However, the surface energies are very close for the FSP AZ31 and AZ31–nHA composite, suggesting minor role of incorporated nHA in improving the surface energy. The specific surface area is increased in a metal as the grain size becomes small [32]. Therefore, fine grain structured materials exhibit higher surface energies (due to availability of higher surface area) than coarse grained materials [33]. From the above observations, it can be understood that the FSP induced grain refinement was the reason behind the increased surface energy in FSPed samples. As the fraction

Element

AZ31 (%)

After FSP (%) FSP AZ31

AZ31–nHA

Mg

96.21

96.10

96.25

Al

2.9

2.78

2.82

Fe

0.002

0.001

0.004

Mn

0.02

0.03

0.02

Zn

0.88

0.82

0.85

Ca

–

–

2.16

P

–

–

1.31

Table 2 shows the elemental composition of the samples before and after processing obtained from ICP-AES studies. It is evident from the results that the dissolution of Fe after FSP was within the tolerance limit and so its effect on the corrosion rate can be expected to almost negligible. Further, from the visual observations, the surface of the tool shoulder and the pin were completely covered by magnesium coating after the processing. This further avoided the dissolution of Fe into the work piece. 3.3 Wettability The static contact angles of AZ31, FSP AZ31 and AZ31– nHA samples are 77.1° (±1.2, N = 5), 64.2° (±2.5, N = 5) and 62.8 (±1.5, N = 5) respectively. The measured contact

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Fig. 4 Typical photographs of the water droplets on the surface of the samples and calculated surface energies

of grain boundary was increased after FSP due to smaller grain size, the excess energy at the grain boundary has led to an increase in the surface energy as calculated for FSP AZ31 and AZ31–nHA samples. 3.4 In vitro bioactivity in supersaturated SBF Figure 5 shows the SEM images of the samples immersed in SBF 59 solution for 24 h. The elemental composition of all the samples obtained using EDS analysis after immersion test are listed in Table 3. Several irregular surface cracks were observed along with some white deposits with wide degraded area on the AZ31 sample (Fig. 5a). From the EDS analysis, the chemical composition was observed as Mg, Al, Zn, O, P, Ca and Cl. The elements Mg, Al, and Zn are originally from the sample while Ca, P and Cl are from the SBF. Presence of Cl on AZ31 is due to formation of MgCl2 as reported by Wang et al. [3]. White precipitates along with Mg(OH)2 in flakes morphology were appeared on the substrate of FSP AZ31 as shown in Fig. 5d. The EDS analysis showing the increased deposition of Ca and P suggests enhanced bioactivity of the FSP AZ31. The morphology of the precipitates on AZ31–nHA composite was different from the other samples (Fig. 5g). These precipitates covered the surface in the form of tiny spherical clusters as shown in the magnified image (Fig. 5h) The EDAX analysis shows Ca and P as the prominent elements. And the Ca/P atomic ratio of the samples is in the order of AZ31–nHA [ FSP AZ31 [ AZ31 (Table 2). The early deposition of more Ca/P mineral phase clearly indicates the excellent bioactivity of FSP AZ31 and AZ31–nHA composite. Moreover, the combined effect of small grain size and incorporated nHA led to rapid

biomineralization in AZ31–nHA composite compared to FSP AZ31. This is due to the increased surface energy and the presence of nHA particles that results in the heterogeneous nucleation and rapid growth of apatite crystals on the surface of the composite. The energy barrier to form homogeneous nuclei over AZ31Mg alloy when it is immersed in SBF can be eliminated totally by providing HA on the substrate. Then the heterogeneous nucleation is initiated, resulting in rapid formation of apatite [34, 35]. Figure 6a shows the SEM morphology of the AZ31 sample immersed for 72 h in SBF 59. The surface was covered with Mg(OH)2 completely. However, from EDS spectrum a little amount of Ca/P mineral phase was observed. On the surface of FSP AZ31 sample, a combination of flake like morphology and precipitates of spherical morphology as shown in magnified image were observed as the immersion time increased to 72 h (Fig. 6b). From the XRD analysis (Fig. 7), the phases on the FSP AZ31 sample were identified as Mg(OH)2 and HA. A thick and dense apatite formation on AZ31–nHA composite can be observed (Fig. 6c) after 72 h of immersion. But, the entire sample surface was not covered with apatite like coating. From the EDS analysis, AZ31–nHA sample has Ca and P as the prominent elements with the atomic ratios in the order of AZ31–nHA [ FSP AZ31 [ AZ31 (Table 2). From the XRD analysis (Fig. 7) of AZ31–nHA sample these phases were confirmed as hydroxyapatite. The peaks (002), (211), (300), (310), (222) and (213) corresponding to HA at 26°, 31.8°, 33°,40°, 46.8° and 49.2° respectively along with the other a-Mg peaks over AZ31–nHA sample after 72 h of immersion indicate the improved biomineralization of the AZ31–nHA composite. The peak intensities were observed to be increased for the composite compared to FSP AZ31 samples after 72 h of immersion. Also, the peaks at 18.6° and 40° corresponding to Mg(OH)2 were observed from the XRD analysis on all the samples. The Mg(OH)2 formation at the early stage of corrosion process acts as a barrier for the further degradation, but the presence of chloride ions contributes to the breakdown of the Mg(OH)2 by forming MgCl2 crystals which increases pit formation. These MgCl2 crystals are easily soluble and cause the degradation rate to increase [3]. From the EDS analysis of the samples it can be observed that the presence of chlorine was more for AZ31 sample immersed for 24 h. For FSP AZ31 and AZ31–nHA composite, as the deposition of Ca and P is increased, presence of chlorine was found to be reduced. The effect was higher in the composite but, less Ca/P atomic ratio observed for all the samples compared to the stoichiometric ratio of HA (1.67) suggest the formation of other phosphate phases as reported by other authors [36, 37]. Deposition of HA along with magnesium phosphate was observed on grain refined pure magnesium immersed in SBF 59 for 72 h in our earlier

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Fig. 5 SEM images of the samples immersed in SBF 59 for 24 h: a AZ31, b corresponding magnified image and c EDS analysis, d FSP AZ31, e corresponding magnified image and f EDS analysis, g AZ31–nHA composite, h corresponding magnified image and i EDS analysis Table 3 Elemental composition of the samples after immersion test observed using EDS analysis Sample AZ31

FSP AZ31

AZ31–nHA

Immersion time (h)

Ca

P

O

Cl

Ca/P ratio

24

0.7 ± 0.1

2.4 ± 0.5

25.5 ± 1.6

51.9 ± 1.2

8.3 ± 1.9

0.29 ± 0.1

48

2.2 ± 1.1

4.5 ± 0.6

23.9 ± 2

58.2 ± 3.8

4.5 ± 4.4

0.48 ± 0.2

72

6.6 ± 1.5

8.9 ± 0.5

11.1 ± 1.2

51.2 ± 1.9

0.6 ± 0.1

0.75 ± 0.1

24

4.2 ± 1.9

5.9 ± 1.7

22.2 ± 3.2

62.9 ± 1.4

0.3 ± 0.2

0.68 ± 0.2

48

6 ± 0.7

8.8 ± 0.8

7.4 ± 1.2

57.7 ± 1.4

0.9 ± 0.2

0.67 ± 0.1

72

11.2 ± 1.7

12.8 ± 1.5

8.9 ± 1.3

59.1 ± 10.1

0.7 ± 1

0.87 ± 0.1

24

8.6 ± 1

11.2 ± 2.3

10.2 ± 2

69.5 ± 4.7

0.7 ± 0.3

0.79 ± 0.2

48

12.7 ± 1.7

13.9 ± 0.4

11.7 ± 4.1

59.9 ± 7.8

0.8 ± 0.5

0.87 ± 0.1

72

16.7 ± 1

15.7 ± 0.5

7.2 ± 0.4

59.7 ± 1.4

0.6 ± 0.5

1.07 ± 0.1

study [38]. However, the corresponding peaks were not observed from the XRD analysis in the present study due to presence of magnesium phosphate in small amounts. The

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Mg

above results indicate that the fine grain structure achieved in AZ31 by FSP and embedded nanoparticles of HA are the two prime reasons behind the enhanced mineralization in

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Fig. 6 SEM observations after 72 h of immersion in SBF 59: a AZ31, b FSP AZ31 and c AZ31–nHA composite (the magnified area of white box shown in the corresponding image)

lower weight loss compared to the other samples. This may be due to the enhanced bioactivity. The pH measurements (Fig. 9b), done for every 12 h, shows that the increase in pH was more for AZ31 due to more release of Mg2? and OH- ions into the solution because of higher degradation rate. The AZ31–nHA composite has marginally higher pH compared to FSP AZ31 as the immersion time increased to 24 h due to the possibility of release of Ca2? ions along with Mg2? and OH- ions from the composite in the course of degradation. However, as the degradation rate was reduced for the composite after 72 h, the pH of both FSPed samples are nearly similar. Moreover, for all the measurement times, FSPed samples have lower pH compared to AZ31.

Fig. 7 XRD patterns of the samples after immersion test conducted in SBF 59

FSP AZ31 and AZ31–nHA composite when immersed in SBF 59 solution. 3.5 Corrosion behavior 3.5.1 Degradation in immersion test Figure 8 shows the typical photographs of the samples before and after immersion test carried in SBF 59 for 72 h. The weight loss after different intervals of time for all the samples is shown in Fig. 9a. The surface of the AZ31 shows rapid degradation including irregular pitting compared to FSP AZ31 and AZ31–nHA composite, whereas, better protection from the aggressive attack of the chloride ions was observed on both the FSPed samples. This is due to the deposition of the mineral phases like HA, magnesium phosphate and the more stable Mg(OH)2 layer on the FSPed specimens compared to AZ31. However, AZ31– nHA composite showed uniform surface degradation and

3.5.2 Electrochemical test The potentiodynamic polarization curves of the samples are shown in Fig. 9c, d and the electro chemical parameters are shown in Table 4. Corrosion potentials (Ecorr) were shifted towards positive side and lower values of current densites (Icorr) for FSP AZ31 and AZ31–nHA compared to AZ31. This indicates the improved corrosion resistance for FSP AZ31 and AZ31–nHA composite before biomineralization study. This behavior after FSP is due to smaller grain size, which helps in developing a quick passive layer due to availability of more grain boundary and the reduced intensity of galvanic couple between grain interior and grain boundary results in uniform corrosion [39–41]. The Ecorr of FSPed samples are very close together; but the Icorr of the composite which is lower than the other samples indicates better corrosion resistance compared to other samples. This can be attributed to combined effect of small size and distribution of HA in the matrix, which reduces the localized corrosion by promoting uniform corrosion [42]. Furthermore, the electrochemical parameters of the samples after 72 h of immersion test indicate the best

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Fig. 8 Typical photographs of the samples before and after 72 h of immersion in SBF 59: a AZ31, b FSP AZ31, c AZ31–nHA before immersion test, d AZ31, e FSP AZ31 and f AZ31–nHA after 72 h of immersion

corrosion resistance in the composite compared to the other samples. The prime reason behind the enhanced corrosion resistance for the composite after 72 h of immersion in SBF is due to the higher deposition of mineral phases as observed in biomineralization studies, which stabilizes the Mg(OH)2 phase and provides a protective layer and reduces the attack of the chloride ions. 3.5.3 Comparison of corrosion rates obtained from electrochemical test and immersion test Comparison of the corrosion rates of the samples calculated from the weight loss after 72 h of immersion and electrochemical test conducted on the samples after 72 h of in vitro bioactivity test are shown in Table 4. The corrosion rate was found to be more for AZ31 from both the tests. Among all the samples, AZ31–nHA composite shows lower corrosion rates in both the immersion and electrochemical tests. Interestingly, the corrosion rates calculated from the weight loss of the samples from the immersion test are more than the corrosion rates obtained from the electrochemical test. Witte et al. [3] has previously

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reported the discrepancy in the corrosion rates obtained by immersion (in vitro) and electrochemical tests for Mg alloys (AZ91D and LAE442). Also, a large amount of difference in corrosion rates measured by in vitro and in vivo was observed. Similar discrepancy between the corrosion rates obtained by weight loss method and electrochemical test was reported for commercial pure Mg, AZ31, AZ80 and AZ91Mg alloys [43] and discrepancy between the corrosion rates obtained by electrochemical test and hydrogen evolution study was also reported [44]. Furthermore, the authors [43, 44] suggested that the anomalous electrochemical behavior of the metallic surface was the prime reason behind the difference in corrosion rates and the effect was more for the commercially pure Mg (ten times) and AZ31 (two times) compared to other AZ series Mg alloys. Also, experimentally, Kirkland et al. [45] explained the possible reasons for the discrepancy in the corrosion rates obtained by immersion test and electrochemical test of Mg–5Ca and Mg–10Zn magnesium alloys. However, considering relative effect rather than absolute value, the corrosion susceptibility of the material can be understood by different methods [43, 45]. In the

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Fig. 9 Corrosion behavior of the sample: a weight loss after 72 h of immersion test, b pH of the solution during the degradation, c potentiodynamic polarization curves of the samples before

immersing the samples in SBF 59 for 72 h and d after 72 h of immersion in SBF 59

Table 4 Electrochemical parameters of the samples from potentiodynamic polarization test and comparison of the corrosion rates obtained from electrochemical test and immersion test Before/after biomineralization study (72 h)

Ecorr (V)

Before

-1.712

FSP AZ31 AZ31–nHA AZ31

Sample

Icorr (A/cm2)

Corrosion rate, CR, (mm/year) Electro-chemical test

Immersion test

15.81 9 10-3

–

–

-1.506

4.67 9 10-3

–

–

-1.564 -1.691

2.314 9 10-4 0.409 9 10-3

– 9.4

– 16.39

FSP AZ31

-1.629

0.318 9 10-3

7.3

8.66

AZ31–nHA

-1.506

0.114 9 10-3

2.62

7.61

AZ31

After

present study, AZ31–nHA composite has relatively lower corrosion rates in both the tests compared to the other samples. Therefore, from the above study it can be understood that the improved corrosion behavior of FSP AZ31 and AZ31–nHA composite is attributed to fine grain structure. Among all the samples, AZ31–nHA composite exhibited lower corrosion rate.

3.6 Cytotoxicity and cell adhesion studies Figure 10 shows the percentage viability of the L6 cells exposed to AZ31, FSP AZ31 and AZ31–nHA for 72 h. All the materials showed negligible toxicity. As the incubation time increased from 24 to 48 and 72 h, cells exhibit similar response to all of the samples. The viability of FSP AZ31

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was marginally reduced as the incubation time increased to 72 h. However, based on the statistical analysis (P [ 0.05), no significant difference between the samples with respect to cell viability was observed. It can be understood from

Fig. 10 Cytotoxicity of the samples expressed as % cell viability using MTT assay (standard polystyrene tissue culture plates were used as control and the percentage cell viability of the samples in each case was normalized to control. Statistically no difference was observed between the samples (P [ 0.05))

the results that there is no adverse affect on the cells after processing due to the small grain size and embedded nHA particles. Figure 11 shows the morphologies of the adhered L6 cells on the sample surfaces. On the AZ31 surface, the cells were grouped and inter connected but adhesion was found to be low when compared with FSPed samples. Whereas, on the surfaces of both the FSP AZ31 and AZ31– nHA it was clearly found that the interconnected cells tend to stretch and adhere to the surface. More number of cells covered the surface of AZ31–nHA and started proliferation. It can be explained that the factors behind the better adhesion of the cells on the AZ31–nHA sample are (i) high surface energy, which characterizes the surfaces with high wettability [46], (ii) embedded nHA particles, which enhances the osseointegration [9] and (iii) lower degradation rate when compared to other samples, which favors the cell adhesion [8]. High wettable surfaces promote adsorption of proteins, which subsequently improves cell adhesion to develop strong bond between material and tissue [31]. If the corrosion rate is high, adhesion and growth of cells are affected by the evolution of hydrogen. It is difficult for cells to proliferate on a rapidly degrading surface. Better corrosion resistance observed for the FSP AZ31 and AZ31–nHA favors the cell adhesion compared to that of

Fig. 11 SEM images of the L6 cells on the samples incubated for 24 h: a standard polystyrene tissue culture plate, b AZ31, c FSP AZ31 and d AZ31–nHA composite

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AZ31 sample. The cells on AZ31–nHA composite seemed to be cover wide surface compared to FSP AZ31 suggesting the combined effect of small grain size and presence of HA favors towards improved cell activities. The above observations demonstrate that the FSP can be a promising technique to develop nHA reinforced AZ31 magnesium alloy to fabricate biodegradable orthopedic implants. Moreover, the process is a solid state technique; the issue concerning the stability of the dispersing phase that is generally associated with melting and sintering methods can be avoided in developing new composites. Also, the processing time required to fabricate the composites compared to other techniques is very low. However, the technique involves few machining operations like groove producing by milling cutter before FSP, cutting and machining after processing to prepare the components with required size for an actual application. Experiments can be designed with multi passes with a single setup to make the process cost effective, increase the productivity and the quantity of nHA dispersed into the matrix as well. Also, as the tool geometry (especially the shoulder diameter, pin profile and length) influences the depth of the effected zone, FSP tool can be designed according to the required thickness of the composite in developing AZ31–HA MMC for biomedical applications, especially targeted for temporary implants like locking plates, bone plates and fixtures of dimensions ranging from 2–4 mm thickness and 10–20 mm width with a desired length.

4 Conclusions Grain refined AZ31–nHA MMC was successfully fabricated by FSP. The nHA particles were found to be dispersed in AZ31 after FSP. Wettability was observed to be increased substantially for the processed surfaces due to the induced fine grain structure. Early formation of mineral phases over the surface of FSP AZ31 and AZ31–nHA composite helped to reduce the localized degradation due to pitting. The AZ31–nHA composite has excellent biomineralization and better corrosion resistance compared to fine grained FSP AZ31 and AZ31 samples due to the combined advantage of small grain size and incorporated nHA particles. All the FSPed samples were biocompatible and cell adhesion was promising for AZ31–nHA composites. Hence, it can be concluded that the FSP can be a new promising approach to develop magnesium based composites for biomedical applications. Acknowledgments Authors would like to thank Prof. K Prasad Rao, IIT Madras for providing the NRB supported FSP facility.

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