A biodegradable polymer-based coating to control the performance of magnesium alloy orthopaedic implants
Introduction
The most commonly used materials for bone fracture fixation are usually made of medical-grade metals such as 316L stainless steel, pure titanium and its alloys, and cobalt–chromium-based alloys [1], [2] which are non-biodegradable. However, one desirable characteristic of an implant is its ability to be degraded after the bone has healed as problems may arise if the implants are not degradable. Long-term adverse effects or even an increased risk of local inflammation may occur after long-term implantation since the metallic implant is a foreign body to human tissues [3]. If this is the case, second surgery is subsequently conducted for implant removal. However, repeated surgery not only increases the morbidity rate of the patients, but also results in an increase of health care costs and longer hospitalization [1]. To reduce such complications, the use of biodegradable metallic implants has been investigated [1], [4], [5], [6], [7].
Magnesium and its alloys are the most commonly used metal amongst all the degradable metallic materials. However, the major obstacles of the clinical use of magnesium-based materials are its rapid degradation rate and the release of hydrogen gas upon degradation [8], [9]. Troitskii and Tsitrin used a magnesium–cadmium alloy to secure various fractures, however, reported that the mechanical integrity of the magnesium alloy was only maintained for 6–8 weeks with the release of hydrogen during the corrosion process [10]. Hence, in order to make use of magnesium-based materials feasible for surgical implantation, the corrosion rate must be controlled.
The enhancement of the corrosion resistance of magnesium can be achieved by using different modification methods such as alloying [11] and various surface treatments [8]. Witte et al. [4], [12] suggested that magnesium alloys, especially those containing rare earth elements seemed to be suitable for use as orthopaedic implants. However, in addition to the alteration of its original mechanical properties, the addition of rare earth metals such as zirconium and cerium into the magnesium substrate may potentially add toxic effects to cells [13], [14], as the cytocompatibility of these elements is not known. In the studies by Li et al. (2008) [5], Zhang et al. (2009) [15] and Zberg et al. (2009) [16]; a magnesium–calcium alloy, magnesium–zinc alloy and magnesium–zinc–calcium alloys, were fabricated respectively and used as biodegradable implants, however, the change in mechanical properties of these alloys during degradation were not addressed.
Apart from alloying, surface modifications such as micro-arc oxidation (MAO), ion implantation and plasma anodisation to improve the corrosion properties of magnesium alloys have been investigated [17], [18], [19], [20], [21], [22]. Electrochemical tests were conducted with those surface-treated samples, and an increase in corrosion resistance was reported [17], [18], [19], [20], [21], [22], however, as the biological integrity of those surface-treated samples was not reported, there was insufficient data to draw any conclusion before applying in clinical use.
In this paper, we improve the properties of magnesium implants via the deposition of a biodegradable polymer-based porous membrane made of polycaprolactone (PCL) and dichloromethane (DCM) onto a commercially available magnesium alloy in order to control its degradation rate. This paper aims at investigating the feasibility of these polymeric membranes in controlling the degradation of magnesium alloy under in-vitro and in-vivo conditions, and addresses the cytocompatibility and mechanical integrity of the deposited samples during degradation.
Section snippets
Sample preparation
An AZ91 magnesium ingot with 9 wt% aluminium and 1 wt% zinc (Jiaozuo City Anxin Magnesium Alloys Scientific Technology Co., Ltd.) was used in this study. Disc samples which were 5 mm in diameter and 4 mm in thickness were prepared for the electrochemical corrosion test, immersion test and in-vitro studies while rod samples were prepared for the mechanical integrity testing and in-vivo animal study. The rod samples for mechanical testing were 3 mm in diameter and 9 mm in length whereas for the
Surface morphology analysis
Fig. 2 shows the surface morphologies of the polymer-deposited samples under scanning electron microscopy (SEM). The total pore area of the LPM sample was found to be approximately 236 μm2 in which most of the pores were between 0.8 μm and 1.6 μm in size. The average pore size was 0.302 μm, and the porosity of the LPM sample was 18.2%. On the other hand, the total pore area of the HPM sample was approximately 572.1 μm2, with the pore size ranging between 3.2 μm and 6.4 μm. The average pore size
Discussion
The use of magnesium alloys as biodegradable materials was first investigated during the first half of the last century [28], [29]. The major obstacles of applying magnesium alloys in clinical use are its rapid degradation rate and the release of hydrogen gas upon degradation. Hence, different modifications of magnesium alloys have been conducted, of which one of the approaches is surface modification [8], [17], [18], [30], [31], [32]. By conducting a suitable surface modification, the
Conclusion
In summary, this study demonstrated the effectiveness of applying a porosity controllable biodegradable polymer membrane on a magnesium alloy. The addition of a polymer-coating on the implant was shown to reduce the corrosion rate of the implant. This was mainly related to the pore size of the membrane, which may be altered during synthesis to suit potential applications. In addition to reducing the corrosion rate of the magnesium alloy, the polymer-coated samples also aided in retaining the
Acknowledgement
This study was financially supported by the Hong Kong Research Grant Council Competitive Earmarked Research Grant (#718507) and HKU University Research Council Seeding Fund.
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The first two authors share the co-first authorship.