Sumit Bahl,
Satyam Suwas and
Kaushik Chatterjee*
Department of Materials Engineering, Indian Institute of Science, Bangalore 560012, India. E-mail: kchatterjee@materials.iisc.ernet.in; Tel: +91-80-22933408
First published on 10th October 2014
Magnesium and its alloys are an emerging class of resorbable materials for orthopedic and cardiovascular applications. The typical strategy underlying the development of these materials involves the control of material processing routes and the addition of alloying elements. Crystallographic texture is known to control bulk mechanical as well as surface properties. However, its role in determining the properties of magnesium for implant materials has not been well studied. In this work, an extruded rod of pure magnesium was cut in multiple directions to generate samples with different textures. It was found that texture significantly affected the strength and ductility of magnesium. Corrosion rates in Hank's solution decreased with the increased presence of low energy basal planes at the surface. In vitro cell studies revealed that changes in texture did not induce cytotoxicity. Thus, the control of texture in magnesium based implants could be used to tailor the mechanical properties and the resorption rates without compromising cytocompatibility. This study elucidates the importance of texture in the use of magnesium as a resorbable biomaterial.
Magnesium and its alloys are an important class of resorbable metallic biomaterials for orthopedic and cardiovascular applications.1,3–6 Magnesium is favored over ceramic and polymeric biomaterials as it offers a good combination of high strength, fracture toughness and elastic modulus, which closely matches that of bone. Magnesium is a part of human metabolism and is found mainly in bone tissues. Therefore, the controlled release of magnesium ions from degrading implants may not be toxic and the ions can be easily excreted out of the body. In fact, the presence of magnesium ions is reported to have a stimulatory effect on regeneration of bone tissue.7
The high corrosion rate of magnesium compared to other metals makes it suitable for resorbable applications. Although this is an advantage, the current magnesium alloys corrode in physiological conditions before the bone is fully regenerated. Implants lose their mechanical strength due to rapid corrosion, leading to their failure.1,8 Therefore, one of the major challenges in the development of magnesium alloys for resorbable biomedical applications is the control of the corrosion rate. The commonly used methods to control the corrosion rate include alloying, processing and surface treatments. Some of the widely studied alloying elements include Ca, Al, Mn, Zn, Zr, and Si.9–12 It has been shown that magnesium alloys processed to obtain fine grain size show an enhanced corrosion resistance.13 Various surface modification techniques such as micro-arc oxidation, alkali treatment, chemical conversion coatings, and sol–gel coatings have been applied to improve the corrosion resistance of magnesium.6,14,15
Apart from corrosion, the other important requirement for an implant is mechanical properties such as strength, ductility and elastic modulus. Incidentally, the corrosion as well as the mechanical properties are strongly dependent on crystallographic texture. Crystallographic texture in a polycrystalline material is defined as an occurrence of preferred orientation of grains within the material. It is widely used in numerous engineering applications to tune material properties. The texture of magnesium alloys is known to significantly affect the corrosion resistance.16,17 Given the anisotropy associated with its hexagonal crystal structure(Fig. 1(a)),18,19 mechanical properties such as strength, ductility, and elastic modulus are especially sensitive to the texture of magnesium. Control of texture to tailor properties has been successfully employed in many industries, especially for automobile and aerospace applications. Although the effect of texture on structural properties and corrosion behavior is well established, its effects, if any, on the biological response to a material are poorly understood. A few recent investigations20–22 have reported effect of texture on the cellular response to titanium and its alloys. However, the use of texture as a tool to tune properties to optimize the performance of materials is yet to be fully exploited for biomedical applications and there are no known reports on resorbable biomaterials such as magnesium.
The aim of this study is to investigate the effect of texture on the mechanical properties, corrosion behavior and biocompatibility of commercially pure magnesium for potential use in orthopedic applications. An extruded rod of pure magnesium was cut in three directions (Fig. 1(b)). Thus, the samples had similar grain size distribution but different textures. The mechanical properties and corrosion resistance for the different textures were systematically characterized. Biocompatibility was evaluated by studying the effect of leached magnesium ions on mouse osteoblasts as a model for their use in orthopedic applications.
An MTT (3-[4,5-dimethylthiazol-2-yl]-2,5-diphenyltetrazolium bromide; thiazolyl blue) assay (Sigma) was used to measure cell viability at 1 day and 3 days after adding the conditioned medium. A stock solution of 5 mg ml−1 was prepared by dissolving MTT powder in sterile phosphate buffered saline (PBS). A working solution was prepared by diluting the stock solution to 1 mg ml−1 in growth medium. 100 μl of working solution was added to each well and incubated at 37 °C and 5% CO2 for 2 h. The dye is reduced to purple colored formazan crystals by dehydrogenase enzyme produced in the mitochondria of viable cells. The crystals were dissolved by adding 150 μl of DMSO to each well and incubating for 15 min at 37 °C. The dye concentration was quantified by measuring absorbance at 570 nm using a micro-well plate reader (Biotek). All data is presented as mean ± S.D. for n = 3.
Mg2+ concentration in conditioned medium was measured using an atomic absorption spectrometer (AAS, Thermo Scientific). A standard curve was drawn using samples containing known concentrations of Mg2+. The conditioned medium was diluted 75 times in ultrapure water (Sartorius) before measurement to bring the concentration within the range of the standard curve. The pH of the conditioned medium was measured using a standard pH meter.
The effects of Mg2+ ion concentration and pH on osteoblast viability were studied. Mg2+ concentration was varied in growth medium by exogenous addition of known quantities of MgCl2. The concentration of Mg2+ in the medium was varied from 102 μg ml−1 to 104 μg ml−1. Growth medium containing an endogenous Mg2+ concentration of 20 μg ml−1 was used as the control. The pH of the medium was varied by adding NaOH solution. The final pH values of the media prepared were adjusted to 8.0, 9.0, 10.0, 11.0 and 12.0. Growth medium of pH 7.4 was used as a control. 200 μl of cell suspension containing 5 × 103 cells was added to each well. The growth medium with controlled Mg2+ concentration and pH was added after 24 h of seeding cells. MTT was used to measure osteoblast viability at 1 day and 3 days after replacing growth medium following the procedures as described above. All data is presented as mean ± S.D for n = 3.
Statistical analyses were performed using one-way ANOVA (analysis of variance) with Tukey's test. Differences were considered significant for p < 0.05.
Table 1 compiles the calculated volume fractions of basal and prismatic fibers for the three samples. The transverse sample has the highest volume fraction of prismatic fiber and lowest fraction of basal fiber, as the prismatic planes are parallel to its surface. The longitudinal sample has the highest fraction of basal fiber and lowest fraction of prismatic fiber as the basal planes are parallel to its surface. The 45° sample has lower volume fractions of both prismatic and basal fiber as its lies in between the prismatic and basal orientations. To summarize the texture of extruded Mg, the prismatic planes are perpendicular and the basal planes are parallel to the extrusion axis as is shown schematically in Fig. 1(b).
Sample | Volume fraction of fibers | ||
---|---|---|---|
(0001) | (10![]() |
(11![]() |
|
Transverse | 0.1 | 18.9 | 19.5 |
Longitudinal | 8.2 | 1.0 | 1.0 |
45° | 0.1 | 5.6 | 6.7 |
Sample | Yield strength (MPa) | Tensile strength (MPa) | Ductility (%) |
---|---|---|---|
Transverse | 46 | 275 | 18 |
Longitudinal | 37 | 211 | 16 |
45° | 43 | 213 | 19 |
The transverse sample was unfavorably oriented for basal slip, but suitably oriented for {102} tensile twinning. However, as the CRSS for {10
2} tensile twinning is higher than basal slip, the transverse sample displays the highest yield strength. The longitudinal samples had an orientation slightly deviated from ideal basal. This orientation can lead to activation of basal slip. Since basal slip has the lowest CRSS among all the deformation modes, longitudinal samples have the lowest yield strength. The 45° sample is also favorably oriented for basal slip and consequently has lower yield strength. The ductilities of the transverse and 45° samples are higher than that of the longitudinal sample. Another important aspect is the nature of the stress–strain curve, which indicates the deformation mechanism operating during the compression test. Transverse samples display low strain hardening rate initially due to twinning. Twinning can provide a maximum compressive strain of 0.065, above which it saturates.25 The results of the present investigation are corroborated by reported literature. A low strain hardening rate can be seen from Fig. 4 at a strain around 0.065, above which it increases. Once the twining saturates, it renders the grain toward hard orientation for basal slip, leading to rapid strain hardening. At higher strains (>0.065) the strain hardening rate increases. The transverse sample initially has a lower strain hardening rate than the longitudinal sample, but increases at higher strains. As a result, the transverse curve superseded the longitudinal curve at a strain above 0.125. A higher strain hardening rate could likely be the reason for the high ductility of the transverse sample. In contrast, a lower strain hardening rate might lead to reduced ductility in the longitudinal sample. The nature of the flow curve of the 45° sample is similar to that of the transverse sample in the initial stages of plastic deformation. Both the curves indicate an initial low work hardening rate, which is likely due to the occurrence of twinning. As the strain increases, the twinning strain saturates and basal slip becomes the dominant deformation mechanism. The strain hardening, therefore, increases and the curve is similar to that of the longitudinal sample at higher strain levels. However, the effect of twinning on strain hardening is less pronounced in the 45° sample than in the transverse sample. The strain hardening rate is higher for the 45° sample than for the transverse sample. Kleiner et al. theoretically calculated the variation of twinning strain as a function of deviation from extrusion direction.19 The twinning strain in compression decreases with deviation from the initial extrusion direction. The lesser propensity of twinning in the 45° sample accounts for the difference in strain hardening rates between the 45° sample and the transverse sample. The strain hardening rate in the 45° sample at higher strain is lower than the transverse sample and similar to the longitudinal sample. Although the total strain is higher in the 45° sample than the longitudinal, the net plastic strain is same as in the longitudinal (∼0.15) sample. Similar values of ultimate tensile strength (UTS) and plastic strain can be attributed to the similar work hardening behaviors of the 45° and longitudinal samples.
Sample | Corrosion potential Ecorr (mV vs. SCE) | Corrosion current density Icorr (mA cm−2) | Basal fiber volume fraction (%) |
---|---|---|---|
Transverse | −1667 | 0.16 | 0.1 |
Longitudinal | −1645 | 0.05 | 8.2 |
45° | −1631 | 0.10 | 0.1 |
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Fig. 6 Absorbance values of MTT assay after 1 day and 3 days. * represents a statistically significant difference (p < 0.05) after 3 days. |
The anodic and catholic reactions that occur during corrosion of magnesium are as follows:
Anodic: Mg → Mg2+ + 2e− |
Cathodic: 2H2O + 2e− → H2 + 2OH− |
Corrosion leads to simultaneous increase in Mg2+ concentration and the pH of the medium due to the formation of OH− ions. The pH values of all the three conditioned media were similar at 8.0 (Table 4). The values of magnesium concentration present in the media conditioned for 24 h measured by AAS are also listed in Table 4. There was no statistically significant difference in concentration among the three samples. Theoretically, the transverse and longitudinal samples should have much higher corrosion rates than the 45°. However, there was no measurable difference in the Mg2+ concentration when the samples were conditioned in growth medium. Mg(OH)2 along with other salts will precipitate following Mg corrosion in the culture medium. All such precipitates were removed by centrifugation and only dissolved Mg2+ ions remained in the conditioned medium for the cell studies. Moreover, an initial high corrosion rate of the transverse sample might have led to the formation of a protective Mg(OH)2 layer, retarding its further corrosion. This does not however mean that Mg2+ concentration is stabilized after 24 h. Mg2+ concentration will increase with time but the rate will be slower due to the presence of the partially protective Mg(OH)2 layer. Conditioned medium was extracted after 24 h to compare the effect of texture on biocompatibility. The conditions are likely to be different in vivo due to the presence of dynamic conditions where the corrosion products will be transported away by biological fluids. Therefore, it is important to note here that Mg2+ concentration in the conditioned medium should not be considered as a true measure of resorption rate. Corrosion rate determined by the Tafel extrapolation method is a more reliable indicator of the resorption rate.
Sample | pH of the conditioned media | Mg2+ concentration in conditioned media (μg ml−1) |
---|---|---|
Transverse | 8.1 ± 0.1 | 766.0 ± 90.0 |
Longitudinal | 8.2 ± 0.1 | 850.0 ± 53.5 |
45° | 8.1 ± 0.1 | 837.5 ± 76.7 |
The results observed for the biocompatibility tests were investigated further. Immersion of magnesium in growth medium leads to aqueous corrosion. It is possible that either increased Mg2+ or pH in conditioned medium affected cell viability. The individual effects of Mg2+ concentration and pH on osteoblast viability were studied. Growth medium was replaced with medium having known Mg2+ concentration and pH. The effect on osteoblast viability was measured after 1 day and 3 days to compare with the effect of conditioned medium on viability after 1 day and 3 days respectively. Fig. 7(a) shows the effect of Mg2+ on cell viability at 1 day and 3 days after replacing growth medium with medium having a known concentration of Mg2+. The concentration was varied from 20 μg ml−1 (growth medium) to 104 μg ml−1. It was observed that after 1 day the cell viability was not affected for Mg2+ concentrations up to 4 × 103 μg ml−1. On the other hand, a concentration of 104 μg ml−1 was found to be toxic as all the cells died within 1 day. Cells were able to proliferate at all concentrations lower than 104 μg ml−1. However, the proliferation rate decreased with an increase in concentration. The cell viability was statistically similar to the control at concentrations up to 750 μg ml−1 but lower at higher concentrations. Thus, concentrations of Mg2+ up to 4 × 103 μg ml−1 are mildly toxic to osteoblasts although the effect is inhibited cell growth. Extremely high concentration (>104 μg ml−1) is toxic to osteoblasts. Hallab et al. found Mg mildly toxic to MG-63 cells at 50 mM (1250 μg ml−1) Mg2+ concentration.29
The effect of pH on cell viability and proliferation is shown in Fig. 7(b). It was observed that after 1 day the cell viability was statistically similar to that of the control (pH = 7.4) up to pH 9.0. The cell viability was significantly lower than the control for pH 10.0, whereas all the cells died at pH levels greater than 10.0. After 3 days the cells proliferated at all pH levels except pH 11.0 and 12.0, where all the cells had died within 1 day. However, the proliferation rate slowed with an increase in pH. The cell viability at 3 days was statistically similar for the control and pH 8.0, but was statistically lower than the control for higher pH values. Thus, these data indicate that although higher pH (<10) inhibits cell growth, it is not cytotoxic. However, extremely high pH (>10) leads to complete cell death. The results are in agreement with the work of Gu et al., where extracts of binary Mg alloys with pH levels around 8.8 did not significantly reduce the viability of MC3T3-E1 osteoblasts.9 From the results of the pH experiment (Fig. 7(b)), it can be said that pH did not play a role in slowing the cell growth in 3 days for all the three samples. The Mg2+ concentration in all three conditioned media extracted after 24 h of sample immersion was approximately 800 μg ml−1. This value is below the limit which can retard osteoblast proliferation (Fig. 7(a)). However, the observed slower proliferation rates in the samples compared to the control could likely be the combined effect of relatively higher pH and Mg2+ concentration. It must be noted that the biocompatibility tests in this study were based on an indirect assay in static culture. The implants in the human body experience a more complex and dynamic environment. Body fluids will transport both precipitated Mg salts as well as dissolved Mg2+ ions away from the implant site. The cytocompatibility will be influenced by the transport of corrosion products with the fluids. Therefore, further studies with animal models are warranted to fully evaluate the role of texture on the biocompatibility of magnesium implants.
Few studies have reported the effect of crystallographic texture on the performance of biomaterials. Faghihi et al.22 showed that texture of Ti-6A-4V alloy affects osteoblast attachment and proliferation through changes in surface energies. In another study, Hoseini et al.21 found that a higher number of Ti–OH bonds are facilitated by presence of close packed planes which indeed improved cell–material interaction. In a recent study, we reported a systematic study on the role of texture in controlling the mechanical properties, corrosion behavior and osteoblast response to titanium.20 Even as magnesium has emerged as a promising candidate for use as a resorbable biomaterial, there is little understanding of the role, if any, of texture on its performance. From the findings of this study, it can be seen that texture influenced mechanical properties and corrosion resistance without a significant change in cytocompatibility. Thus, texture can be utilized to tune the mechanical properties and resorption rates without the loss of biocompatibility while designing magnesium-based resorbable implants, in addition to other strategies such as alloying, processing and surface modification.
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