Biodegradable magnesium-based biomaterials: An overview of challenges and opportunities

Abstract

As promising biodegradable materials with nontoxic degradation products, magnesium (Mg) and its alloys have received more and more attention in the biomedical field very recently. Having excellent biocompatibility and unique mechanical properties, magnesium-based alloys currently cover a broad range of applications in the biomedical field. The use of Mg-based biomedical devices eliminates the need for biomaterial removal surgery after the healing process and reduces adverse effects induced by the implantation of permanent biomaterials. However, the high corrosion rate of Mg-based implants leads to unexpected degradation, structural failure, hydrogen evolution, alkalization, and cytotoxicity. To overcome these limitations, alloying Mg with suitable alloying elements and surface treatment come highly recommended. In this area, open questions remain on the behavior of Mg-based biomaterials in the human body and the effects of different factors that have resulted in these challenges. In addition to that, many techniques are yet to be verified to turn these challenges into opportunities. Accordingly, this article aims to review major challenges and opportunities for Mg-based biomaterials to minimize the challenges for the development of novel biomaterials made of Mg and its alloys.

Keywords: biodegradability; biomaterials; magnesium; medical device; tissue engineering.

FIGURE 1

In vivo angiography and the follow‐up IVUS results. The in vivo aortic angiography indicating no in‐stent restenosis and late thrombogenesis in the JDBM‐2 alloy and 316L stainless steel (SS) (served as negative control) vascular stents after implanting in rabbit abdominal aorta for 1, 2, 4, and 6 months. The corresponding follow‐up IVUS photographs demonstrating the longitudinal reconstruction of the abdominal aorta after the JDBM‐2 alloy and 316L SS stenting. Enhanced vessel size and lumen patency at various implantation time with the absence of neointimal hyperplasia show that JDBM‐2 alloy can be a promising alloy for vascular stent application. Adopted with permission

FIGURE 2

The degradation process of Mg biomaterials. (A) Micro‐CT photographs of the implanted high‐purity magnesium screw in goats. (B) The SEM morphologies of the screws after 4, 8, 12, and 48 weeks implantation. (C) Quantification of the degradation of high‐purity Mg screws as evaluated by the reduction in surface area of the screws at the given time points, showing that the main body of the screws and the threads were significantly degraded at a period of 48‐week after surgery. At this time of implantation, the femoral head of the goat with the expected shape was healed. Overall, high‐purity Mg screws indicated adequate mechanical integrity and satisfactory degradation kinetics compatible with the healing procedure. (D) Macroscopic display of the FAsorbMg staples while immersing in the artificial intestinal juice and (E) when staples were removed from the artificial intestinal juice at 1, 4, 8, and 12 weeks. A white layer of corrosion products formed on the surfaces of the staples. H₂ gases produced from the fast corrosion of Mg were formed in the first week of immersion assay only in trace amounts, and most staples kept their shapes and mechanical integrity until at least the fourth week. In vitro results showed satisfactory biodegradation behavior, mechanical durability, and biocompatibility of Mg alloy staples. (A‐C) Adopted with permission. (D and E) Adopted with permission

FIGURE 3

Schematic illustration of the corrosion behavior of biodegradable magnesium biomaterial in physiological conditions and several possible chemical reactions. (A) The corrosion of Mg results in the production of Mg²⁺ cations, H₂ bubbles, and OH^‒ ions. (B) A white compound with low solubility in water chemically named Mg(OH)₂ forms on the surface of Mg. (C) Chloride ions attack the surface of Mg and disrupt the protective layers on that, resulting in the formation of MgCl₂ and a higher corrosion rate. (D) OH^‒ ions react with bicarbonate, thus producing carbonate ions and water. (E) MgCO₃ formation owing to the presence of Mg²⁺ and carbonate ions. (F) HPO₄ ^2‒ reacts with OH^‒ ions and generates phosphate ions and H₂O. (G) Phosphate ions react with Mg²⁺ cations, thereby producing Mg₃(PO₄)₂. (H) Calcium carbonate is formed due to the presence of Ca²⁺ and carbonate ions. (I) Calcium phosphate biomineralization, because of the presence of calcium and phosphate. Several possible corrosion products here are tricalcium phosphate and hydroxyapatite (HA), besides X can be, OH, Cl, and so forth. (J) Adherence of cells and proteins to the surface, thereby slowing down the corrosion rate. (K) The hydroxide ions might harm cells. (L) An increase of hydroxide ions leads to an antibacterial influence. (M) As the level of OH^– increases, the more basic, or alkaline, the physiological solution becomes, which decelerates the corrosion rate of Mg. (N) The adsorption of proteins on the surface of the Mg implant is followed by cellular attachment, migration, proliferation, and complex Mg²⁺ ions. (Source: Authors)

FIGURE 4

High‐purity magnesium screws for the fixation of femoral fracture and its implantation in the unscathed femoral. (A) Diagrammatic drawing of a rabbit displaying the high‐purity magnesium screws for fixing the femoral fracture (right leg of the rabbit) and the implantation of the same screw as a control sample in the (left leg of the rabbit). In the right leg, the gray part shows the fracture space of three millimeters. (B) A high‐purity Mg screw before surgery. (C) In vivo three‐dimensional photographs and (D), μ‐CT scans of high‐purity Mg screws both in the fracture and the control groups at 4, 8, and 16 weeks postimplantation. In the fracture group at 4 weeks, pure Mg screws had a bending angle at the part subjected to the fracture because of the mechanical stresses involved in fracture fixation. At 8 weeks postoperation, new tissue was formed in the gap. At 16 weeks, the formed bony bridge was developed, pure Mg screws were severely corroded, and mechanical integrity was reduced. Red arrowheads display the fracture gap. Adopted with permission

FIGURE 5

A comparative image of hydrogen evolution behavior of three different Mg alloys during corrosion. Evaluating hydrogen gases accumulated in the body of mice through implanting three Mg alloy (AZ31, WKX41, and ZJ41) discs with different corrosion resistance in mice by (A) a hydrogen microsensor. A week after surgery, the results indicated that (B) AZ31 Mg alloy with the highest corrosion resistance produced H₂ gas too slowly to show a visible gas cavity. (A and B) Adopted with permission. (C) WKX41 Mg alloy with moderate corrosion resistance generated a smaller H₂ cavity in comparison with (D) ZJ41 Mg alloy with high corrosion rate, which formed a big gas cavity. (E) H₂ concentration on each point from a calibration curve demonstrated that among these three alloys, the ZJ41 alloy presented the highest corrosion rate. (F) Hydrogen concentration was evaluated weekly over a 4‐week study, which indicated that the maximum H₂ concentration produced by these alloys and the percentage decreased in hydrogen both declined in the order of ZJ41 > WKX41 > AZ31. C‐F Adopted with permission

FIGURE 6

Surface modifications of Mg alloys for biomedical applications. (A) Fluorescence photographs of MC3T3‐E1 preosteoblasts cultured on the surface of untreated and (B) PIII‐treated ZK60 alloy substrates for 1 and 3 days indicating that although both untreated and treated ZK60 Mg alloys were biocompatible, the higher MC3T3‐E1 preosteoblasts adhesion occurred on the PIII‐treated ZK60. (C) The fold change of BrdU incorporation after culturing MC3T3‐E1 preosteoblasts for 1 and 3 days with untreated and PIII‐treated ZK60 substrates demonstrated that the amount of incorporation of BrdU in the PIII‐treated ZK60 was 1.7‐fold and 2.5‐fold higher at 1 and 3 days, respectively, showing better osteoblasts viability and proliferation due to the controlled release of magnesium ions from treated ZK60 alloy. Adopted with permission. (D) Schematic of coatings on the surface of Mg and its alloys. (E) Schematic demonstrations of degradation mechanism of MAO/PLLA composite coatings on the surface of Mg–1Li–1Ca alloys in Hanks' Balanced Salt Solution (HBSS) indicating the corrosion of the substrate at the initial stage and swelling of PLLA polymer on the surface of Mg alloy and (F) the final peeling‐off of the PLLA coating on the surface of the alloy under the pressure of corrosion products and hydrogen gas. Adopted with permission