Corrosion behaviour of AZ31 magnesium alloy with different grain sizes in simulated biological fluids☆
Introduction
Magnesium is present in high concentrations in sea water and is the eighth most abundant element on Earth. It has also excellent specific strength and low density, only two-thirds that of Al, so Mg and its alloys can be used in many applications including computer parts, mobile phones, aerospace components, handheld tools, etc. [1], [2]. Mg alloys are also potentially useful for bone implants and stent applications due to their low density, inherent biocompatibility and adequate mechanical properties, including a fracture toughness higher than that of ceramics [3], [4], [5], [6]. Additionally, the elastic modulus of Mg alloys (40–45 GPa) is closer to that of human bones (10–40 GPa) than other commonly used implant materials. As a result, the stress-shielding phenomena caused by current metallic implants made of stainless steel or Ti alloy can be minimized [7].
Another advantage of Mg in relation to other metallic implants is the degradability of Mg alloys which offers the possibility of better physiological repair and better reconstruction of vascular compliance with minimum inflammatory response [8]. It has been shown that magnesium apatite precipitates on the surface of the modified pure Mg [9], [10] and osteoblasts respond by degrading Mg alloys in guinea pig femur [11]. Mg alloys have also shown osteoconductive and osteoinductive properties, thus offering a less invasive repair and temporary support during tissue recovery [7]. The alloys are gradually dissolved or absorbed by the body. In this regard, they are superior to permanent implants, which may cause mismatches in behaviour between the implant and the body as well as physical irritation and chronic inflammatory reactions. As degradable materials, they will not remain as permanent implants in the body and will not require a second surgical operation after the tissue is repaired. Evaluation of the cytotoxicity of Mg and its alloys is reported in Ref. [12].
The major drawback of Mg alloys is that they tend to corrode very quickly in the physiological pH (7.4–7.6) environment, thereby losing their mechanical integrity before the end of the period necessary for bone tissue healing. Therefore, the promising future of Mg and its alloys is dependent on being able to control the rate of corrosion in body fluids.
Several treatments to protect Mg against corrosion have been proposed, such as Mg purification [13], fluoride conversion coatings [14], alloying with other elements and anodizing [1], [13], [15]. Several studies [16], [17], [18], [19], [20], [21], [22] have shown that the corrosion behaviour of Mg alloys is significantly dependent on the microstructure and particularly on the amount and distribution of the intermetallic phases and the grain size. Attempts at improving the corrosion resistance of Mg alloys by their reducing grain size have been proposed by means of laser fusion technology. However, the results obtained by various authors are contradictory [23], [24]. Microcrystallization has been also proposed as a way to improve resistance in chloride media [25].
It can therefore be concluded that the reduction of the corrosion rate, by grain size refinement, and the changes in this rate as a function of the physiological media is not well established in the case of Mg alloys. The aim of this work is to evaluate these two effects in an attempt to develop a Mg alloy with improved corrosion resistance.
Section snippets
Materials
The AZ31 alloy was received from Magnesium Elektron in the form of a rolled 3 mm thick sheet in the O-temper condition (annealed at 345 °C) and also in the form of a cast ingot. The chemical composition of both types of AZ31 alloy was determined by wavelength dispersion X-ray fluorescence (WDXRF) to be: 3.41 ± 0.09 wt.% Al, 0.841 ± 0.039 wt.% Zn, 0.176 ± 0.013 wt.% Mn (balance Mg). The alloy was solution heat-treated at 450 °C for 30 min and water quenched. Rolled sheet disks of 10 mm in diameter were spark
Microstructural characterization
The microstructures of type I and type II AZ31 alloy are shown in Fig. 1a and b, respectively. True grain sizes were measured by the linear intercept method using a correction factor of 1.74. The mean grain sizes were d = 25.7 and 4.5 μm for type I and type II samples, respectively. In both type I and type II samples, some particles of Mn-rich phase (Al6Mn) still remained in the microstructure after processing. Their volume fraction was estimated by image analysis to be close to 3 × 10−3.
Initial stages of corrosion
Discussion
The application of Mg and its alloys as biomaterial for temporary implants in the form of plates and screws would be effective if the corrosion kinetics could be understood and controlled in body fluids. A possible route to reduce the corrosion rate could be the reduction of the grain size. With this goal in sight, the research developed in this work consisted mainly of the evaluation of the corrosion behaviour of type I and type II AZ31 Mg alloy of different grain sizes in two electrolyte
Conclusions
- 1.
At initial stages, pits are associated with the presence of AlMn intermetallics. The monophasic AZ31 alloy shows pitting corrosion that spreads laterally for both types of samples.
- 2.
For short periods, the initiation of localized corrosion occurs earlier in PBS than in NaCl. However, the precipitation of P-containing salts in PBS decreases the corrosion rate over time for both types of samples.
- 3.
The best corrosion behaviour corresponds to the AZ31 alloy with the finest grain size in PBS. It reveals
Acknowledgements
The authors acknowledge the Ministerio de Educación y Ciencia, Spain, for financial support by projects MAT 2006-02672 and MAT2008-06719-C03-01.
D.P. and M.F.L.M. acknowledge the financial support by ANPCyT (PICT 05-33225, PICT 05-32906), UNLP (11/I129) and CONICET (PIP 6075).
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Part of Thermec’2009 Biodegradable Metals Special Issue, edited by Professor Diego Mantovani and Professor Frank Witte.