Mg-Li-Based Alloys as Implant Materials

Abstract

This chapter is aimed at discussing the prospect of using novel magnesium-lithium-based alloys for temporary implantation. It discusses the challenges of implant materials and focuses on the design, characterization, and assessment of Mg-Li-Zn-Ca alloys. Biodegradable magnesium alloys have recently been the material of choice for the manufacture of implantable medical devices because they proffer efficacious solutions to temporary implantation. Magnesium-lithium-based alloys are a unique system of alloys that offer enhanced ductility and uniform degradation. The increase of lithium in the quaternary Mg-Li-Zn-Ca system resulted in phase transformation of the hcp crystal structure of magnesium to bcc, thus improving ductility. Lithium promoted the formation of a solid solution and a compact surface oxide that decreased corrosion kinetics in biological media. The alloys exhibited good biocompatibility, as evidenced by cell viability and metabolic activity when exposed to solutions retrieved from immersion tests. Furthermore, the improvement in mechanical properties and degradation properties of these alloys relative to other magnesium-based alloys provide an opportunity for wider adoption in the biomedical field.

1. Introduction

In the last two decades, there has been a paradigm shift aimed at developing biomaterials for temporary implant applications such as screws and vascular stents. This is because it is not beneficial for metallic materials that have finished serving their purpose to remain in the body permanently. Complications associated with the use of permanent implants for temporary treatments include extractive surgeries, thrombosis, extended use of antiplatelet drugs, and blood thinners. In addition, there are pediatric cases where children are deemed unfit for permanent implant placement. Bioresorbable alloys are required to manage orthopedic, cardiovascular, and neural diseases.

Historically, metals and other materials have been used to repair the human body, dating back several millennia [1]. Until the mid-nineteenth century, copper and bronze were suitable implantation materials, but poisoning from copper ion leaching was a concern [2]. In 1880, Gluck, used ivory prosthesis as implants in the body, and in 1902, gold was used as the interphase between the articular heads of the implant. This experiment proved to be successful, which led to further studies on chemically inert and stable materials [3].

Biodegradable materials dissolve within the human body after serving therapeutic functions, and the healing process is complete. A bioresorbable vascular stent, for instance, should be able to provide the required mechanical strength for restructuring a diseased artery over a specified period and subsequently be gradually resorbed by the body after its function is completed. Bioresorbable materials should be biocompatible, functional, durable, and safe before being considered for implantation. These characteristics determine the type of material that should be used for specific applications. For example, metallic materials, when compared to polymeric materials, have better desirable mechanical properties and are radiopaque [4], thus making imaging easier. Table 1 highlights the essential properties of magnesium alloys used for the manufacture of stents.

Aspect	Description
Resorption	Mechanical integrity 3 to 6 months; Full dissolution within 1 to 2 years
Biocompatibility	Non-toxic, no inflammatory tissue response; No harmful release and/or residue of particles
Mechanical Properties	Tensile yield stress TYS > 200 MPa; Ultimate strength UTS > 300 MPa; Tensile elongation >15–18%
Microstructure	Maximum grain size of 10–12.5 μm
Hydrogen Evolution	Evolution <10 μL H^2 cm^−2 day^−1
Corrosion Rate	Corrosion rate < 0.2 mm/year

Table 1.

Design parameter for biodegradable magnesium stent [5, 6, 7].

Magnesium is a good fit for biological implants because the human body can innocuously process a relatively high level of magnesium content. Magnesium serves multiple biological functions within the body, such as regulating muscles, heart rhythm, cholesterol production, and blood pressure. Moreover, more than 300 enzyme processes within the body require magnesium [8], and it has a recommended daily average intake of 310–420 mg [9] compared to zinc (8–40 mg/day) and iron (5–27 mg/day) [10]. However, the greatest challenge facing the adoption of magnesium-based alloy in the medical industry is the need for controlled and uniform degradation, as well as sustained mechanical integrity [11]. The degradation issues arise from the inherent vigorous electrochemical reactivity of Mg alloys. Magnesium is found to react spontaneously in aqueous solutions because of the ease of transfer of its electrons. Research efforts have considered the use of various alloying elements, heat treatments, and metallurgical processing techniques, such as extrusion, to alter their microstructure, which ultimately impacts degradation and mechanical properties.

The global bioimplant market size is currently estimated at over $260 billion, with increasing demand for minimally invasive surgeries, chronic disorders, bone degeneration, and a rising aging population as driving forces [12, 13, 14]. Orthopedic and cardiovascular implants account for ~50% of the market share by type, whereas metallic implants have the largest share by materials, as shown in Figure 1 [13, 15]. The regional outlook puts North America ahead of Europe and Asia as having the highest market share, with Asia having the fastest-growing market.

Figure 1.

Bioimplants market share by (a) type of implant, (b) implant material, and (c) region [13, 14, 15].

The metallic and medical alloy industry is mostly dominated by permanent implants such as titanium, cobalt chromium, and stainless steel. However, emerging biodegradable materials like Mg-based alloys compete with orthopedic implants and cardiovascular stents. They serve as improved substitutes for previously used biodegradable polymeric devices due to their load-bearing capacity [16]. Other biodegradable metallic materials such as Zn- and Fe-based alloys have not been as successful as their Mg counterpart in application. Mg alloys are currently being developed for various biomedical applications and degrade at different rates to match specific needs.

2. Microstructural and mechanical properties

Magnesium has a hexagonal closed pack (hcp) crystal structure, which is beneficial for strength but detrimental to ductility. This is because an hcp crystal structure does not possess enough independent slip systems required for dislocation movement and uniform deformation. Introducing other alloying elements into magnesium can alter its texture and improve mechanical behavior. Several magnesium-based alloys with biodegradable and biocompatible alloying elements have been investigated for biomedical applications. Most of the studies are focused on improving mechanical performance and combating high initial degradation rates, localized degradation, and hydrogen evolution. For instance, addition of rare earth (RE) metals or manganese (Mn) improves strength, zirconia (Zr) provides grain refinement and increases corrosion resistance, zinc (Zn) also increases corrosion resistance and strength, whereas calcium (Ca) reduces oxidation and allow for easy rollability. Lithium (Li) is one element that stands out because it transforms the hcp crystal structure of magnesium to a body-centered cubic (bcc), thereby increasing the ductility of the alloy and, in some cases, making it superplastic. In addition, Li also provides very good corrosion resistance by developing a uniform oxide surface coating. Figure 2 shows a plot of the ultimate tensile strength and elongation of some researched Mg-Li-based alloys. Careful introduction of metallic elements, both the choice of elements and the alloying percentages, in a fashion that will enhance both mechanical and degradation performance is of optimum importance.

Figure 2.

Ultimate tensile strength vs. elongation for select Mg-Li-based alloys.

When lithium is added to magnesium, it forms a solid solution of α-Mg up to ~5.3 wt.% Li, after which dual phase α-Mg and β-Li coexist in equilibrium up to 10.7 wt.% Li. Further increase in Li content results in phase transformation into a single β-Li phase, as shown in the Mg-Li binary phase diagram of Figure 3. These changes in crystal structure yield corresponding improvements in room temperature ductility but can negatively impact mechanical strength.

Figure 3.

Equilibrium binary phase diagram of magnesium and lithium [17].

2.1 Crystalline phases and phase transformations

The crystalline phases present in an alloy makeup its microstructure; they determine crystal structures, textures, and deformation behavior. In addition, phase transformations such as those induced by the addition of Li to Mg also impact the mechanical and corrosion properties of the alloy. In a study of Mg-xLi-1Zn-0.5Ca (x = 0, 4, 8, 11) cast alloys [18], x-ray diffractograms shown in Figure 4 revealed that the predominant peaks for alloys with 0 and 4 wt.% Li (L0 and L4) were those of magnesium. Minor diffraction peaks of Mg₂Ca and Ca₂Mg₆Zn₃ phases were also detected. For the alloy with 8 wt.% Li (L8), peaks of both magnesium and lithium phases, as well as the secondary Mg₂Ca phase, were detected, whereas the alloy with 11 wt.% Li (L11) displayed a predominant lithium phase with a secondary Mg₂Ca phase. The primary crystalline phase of both alloys L0 and L4 had the hcp structure; that of L8 had a combination of hcp and bcc while L11 had only bcc. The primary crystalline phases of these alloys agree with those of the Mg-Li binary phase diagram.

Figure 4.

X-ray diffractograms of Mg-xLi-1Zn-0.5Ca alloys.

2.2 Strength and ductility

Strength and ductility are two fundamental properties that are assessed for biomedical alloys. A tensile or compressive test can be used to determine the yield and ultimate strengths, and elongation of alloys. Mechanical property requirements may vary for specific implant applications, but alloys are usually expected to have a minimum strength and ductility to enable easy processing and fabrication of devices. Other mechanical tests that could be assessed include hardness, strain hardening, fatigue strength, and wear resistance.

Tensile test data shown in Figure 5 shows how Li content affects the strength and ductility of alloys, as reported in Section 2.1. Increasing the Li content progressively increases the ductility of the alloy across the three phases, but the strength is maximum in the dual phases. Although strength can be attributed to several factors such as grain size, presence, and distribution of secondary phases, here, the alloying content and proportion of both the α-Mg and β-Li phases also play a role. Several thermomechanical techniques such as extrusion, equal channel angular pressing (ECAP), and friction stir process (FSP) improve the strength and ductility of Mg alloys, but Mg-Li alloys tend to possess very good elastic and superplasticity, especially in the duplex phase [19].

Figure 5.

Stress–strain curves and tensile parameters of Mg-xLi-1Zn-0.5Ca alloys.

The mechanical properties of biodegradable implants can be tuned by alloying, thermomechanical processing, surface modification, polymer coating, and other such treatments. The initial mechanical properties should account for the effect of biodegradation, such that the implant devices are able to provide the required mechanical support throughout the healing time.

Figure 6 demonstrates the changes in mechanical strength with degradation time for the L8 and L11 alloys. Specimens of these alloys were subject to accelerated durability tests in saline solution via pulsatile loading with the average regular human heart rate of 72 beats per minute (bpm) for a specified time. The average stress-strain plots after compression tests revealed a higher maximum stress of ~228 MPa for the L8 alloy compared to the L11 alloy (~158 MPa) before durability testing. In contrast, the L11 possessed a higher maximum strength of ~79 MPa compared to L8 (57 MPa) after pulsatile loading. This represents about 50% and 75% loss in strength for alloys of L8 and L11, respectively, after about 3 months implantation period.

Figure 6.

Stress–strain curve of Mg-8Li-1Zn-0.5Ca (L8) and Mg-11Li-1Zn-0.5Ca (L11) before and after accelerated durability tests.

3. Electrochemical behavior

Generally, corrosion is unwanted in engineering and science applications. However, in the case of biodegradable implants, controlled degradation can revolutionize biomedical implantation. The biodegradation of materials refers to the gradual removal and atomic disintegration of materials in a biological environment. This process is characterized by electrochemical reactions and can be studied by the corrosion processes of metals. Corrosion normally occurs at a rate determined by equilibrium between opposing electrochemical reactions. An anodic reaction occurs when the metal is oxidized, releasing electrons into the metal. A cathodic reaction is in which a solution species (often O₂ or H⁺) consumes the electrons released from the metal. When these two reactions are in equilibrium, the flow of electrons from each reaction is balanced, and no net electron flow (electrical current) occurs. The two reactions can take place on one metal or on two dissimilar metals (or metal sites) that are electrically connected.

The composition and strength of the electrolyte or environment in which the material is housed plays an important role in its degradation behavior. Degradation behavior can be assessed in vivo, where the material under test is implanted within a living organism, or in vitro, where the materials are tested in simulated biological conditions. In vivo testing provides more accurate conditions for assessing implant materials but does not allow the variation of single parameters within the study.

Mg alloys degrade in aqueous environment to produce Mg hydroxide (Mg(OH)₂) and hydrogen gas (H₂) with an overall electrochemical reaction shown in Eq. (1). When chloride concentration rises above 30 mmol/l, Mg hydroxide starts to convert into highly soluble Mg chloride [20]. Low hydrogen over potential enables hydrogen evolution and galvanic corrosion, especially when secondary phases are present. It is important to note that although precipitation strengthening and grain refinement improve strength and ductility, care should be taken to prevent localized degradation resulting from galvanic pairs:

$$Mg+\text{2}{H}_{2}O\to Mg{\text{(}OH\text{)}}_2+H_2$$E1

Implant degradation is dependent on several factors such as material composition, device geometry, and environment. The rate of electrochemical reaction and formation and stability of oxide films also depend on reduction and oxidation reactions, time of exposure, and temperature. The corrosive environment of the human body is dependent on the composition of blood and other body fluids, as well as the body temperature. Furthermore, the effect of protein adsorption depends on metal/biomolecule combinations and can promote or reduce degradation rates for different cases [21].

3.1 Factors affecting corrosion behavior

The most important requirement of a biodegradable implant is – its degradation properties in a specific biological environment, which are influenced by various intrinsic and extrinsic factors. Intrinsic factors include chemical composition, microstructure, surface energy, wettability, thickness, and stability of the passivating oxide film. Extrinsic factors include temperature, pH, dissolved oxygen content, amino acids, biomolecules, and chloride ions in the surrounding environment [22]. Tissue fluids within the human body present a very corrosive environment for implant devices. In addition to inorganic species in body fluid, different types of biomolecules and cells may attach to the implant surface and affect its surface chemistry. The regeneration of a passivating oxide film is delayed since the concentration of dissolved oxygen in body fluids is approximately one-fourth of that in air. Additionally, the concentration of Cl⁻ ions in serum and interstitial fluid is 113 and 117 mEq/L, respectively [22]. Cl⁻ ions induce pitting corrosion at sites where the passive film is broken. This is followed by the propagation of the pit, at a rate that occasionally increases with time, because of the increasing acidity inside the pit.

3.1.1 Corrosion kinetics

Electrochemical reactions have been identified to proceed in a linear, parabolic, or logarithmic fashion [2]. The parabolic rate law is attributed to the diffusion of solvent ions through a porous film produced during the corrosion reaction, thus retarding its rate. The rate-determining step is diffusion through a passivating film; thickness of film increases in proportion to the extent of corrosion. For linear oxidation, the rate of oxidation or corrosion generally depends on the pore volume and the “tortuosity” of the pores as well as the film thickness. This is because highly porous, poorly adherent, and fractured non-protective oxide layers do not retard ionic diffusion and the rate of corrosion.

3.1.2 Oxide formation and pilling-Bedworth ratio

Electrochemical redox reactions involve the transfer of electrons and change in oxidation states of elements present within the alloy. Oxide formation is constrained by reaction thermodynamics with a driving force given by the potential difference of redox half-reactions. Oxides with minimum energy of formation are preferentially formed, and a lower enthalpy of formation is indicative of greater chemical stability [23]. However, the effective driving force of metal oxidation reduces as the oxide grows because of potential distribution [24]. The Pilling–Bedworth ratio (PBR) is used to describe the volume ratio and stress states of metals and their oxide films. Generally, a ratio less than 1 is indicative of a thin oxide coating with tensile stresses, which offers poor protection, and a ratio greater than 1 indicates the oxides are denser with compressive stresses and offer more protection. Higher ratios indicate growth in stress, and beyond 2, large compressive stresses result in cracking and spallation of the oxide film. The PBR of compounds can be calculated using Eq. (2):

$$PBR=\frac{M_{oxide}\times {\rho }_{metal}}{n\times M_{metal}\times {\rho }_{oxide}}$$E2

where M and ρ are the respective molecular weights and densities, and n is the number of metal atoms in the oxide molecule (Table 2).

Compound	MgO	CaO	ZnO	Li2O	LiOH	Mg(OH)2	MgCO3	CaCO3	Li2CO3
PBR	0.80	0.64	1.59	0.57	1.26	1.80	2.04	1.43	1.35

Table 2.

Pilling–Bedworth ratios of select compounds.

3.2 Electrochemical methods for accessing corrosion behavior

Electrochemical techniques are ideal for the study of corrosion processes because they provide accelerated corrosion rates as opposed to the conventional weight loss/gain method that requires an extended period for measurements. In electrochemical studies, a metal sample of a known surface area is used to model a redox reaction occurring on the surface of a metal immersed in an electrolyte. The potential between the metal and a reference electrode is varied using a potentiostat, and the current flowing through a counter electrode is measured as a function of potential. The corrosion rate is governed by Faraday’s law given in Eq. (3):

$$Q=\frac{nFW}{M}$$E3

where Q is the total current (coulombs), n is the number of electrons involved in the electrochemical reaction, W is the weight of the metal (grams), M is the molecular weight (grams), and F is the Faraday’s constant (96,485 coulombs/mole).

Accelerated corrosion tests can be used to assess the bio-electrochemical, electro-physicochemical, and electrochemical degradation of Mg alloys within a specific environment. Usually, a potentiostat connected to a three-electrode setup with a standard, counter, and working electrode (the alloy under test) is used to measure the corrosion rate. Different types of accelerated corrosion tests are available to study the degradation behavior of bioresorbable alloys such as potentiodynamic polarization, linear polarization, cyclic voltammetry, and electrochemical impedance spectroscopy.

3.2.1 Potentiodynamic polarization

Potentiodynamic polarization (PP) uses a wide range DC potential to scan the alloy, causing redox reactions and generating corrosion current. The polarization curve can be used to determine the corrosion rate of the alloys from the Tafel slope and according to Faraday’s law in Eq. (4). This technique requires the corrosion potential to remain the same during the measurement to ensure that the applied overvoltage is known. When using the open circuit voltage (Eoc) for polarization, it is imperative to allow sufficient time for the electrochemical double layer to achieve a steady state so that Eoc is stabilized:

$$C.R.=\frac{I_{corr}\times K\times E.W}{A\times \rho }$$E4

where K is a constant, Icorr is the corrosion current, E.W is the equivalent weight, A is the exposed area, and r is the density of the alloy.

3.2.2 Electrochemical impedance spectroscopy

Electrochemical impedance spectroscopy (EIS) measures the response of the alloy under test to AC perturbation. In this technique, it is assumed that the following conditions are met.

1. The response of the system to external excitation can be described by linear differential equations.

2. The system is stable and returns to its previous state after the removal of the external excitation.

3. There is no response before the excitation.

4. The response to excitation has a relationship that is finite.

For an applied signal, Et, which depends on time and frequency, there is a response signal. According to Ohm’s law, an impedance, Z, can be calculated and represented in terms of magnitude and phase shift angle. The measured data resulting from the development of a potential difference between the electric double layer is modeled to an equivalent electrical circuit.

In EIS, the impedance of the corroding metal (working electrode) due to an applied sinusoidal potential change (AC voltage) is analyzed as a function of frequency ω. At each frequency, the resulting sinusoidal current waveform and the applied potential are out-of-phase by phase angle (θ), whereas the current amplitude is inversely proportional to the impedance of the interface. The electrochemical impedance, Z (ω), is the frequency-dependent proportionality factor in the relationship between the voltage signal and the current response, as given in Eq. (5)

$$\text{Z}\left(\omega \right)=\text{E}\left(\omega \right)/\text{i}\left(\omega \right).$$E5

where E is the voltage signal, E = E0 sin (ω t); i is the current density, i = i0 sin (ω t + θ); Z is the impedance (ohm/cm²); and t is the time (seconds).

The impedance is a complex number described by the frequency-dependent modulus, |Z|, and the phase angle, θ, or, otherwise, by the real and imaginary components, Z′, and Z′′ [25]. In electrochemical impedance analysis, three different types of plots are commonly used, including two bode plots, which show impedance and phase angle against frequency. The third is a Nyquist plot, which shows complex plane Z′′ vs. Z′, and the capacitive arc provides an estimate of the corrosion resistance of the material. The relative diameter of the arc is directly proportional to the charge transfer resistance or polarization resistance (Rp). Thus, an increase in semicircular diameter corresponds to an increase in corrosion resistance.

It should be noted that since real electrochemical processes hardly show pure capacitance, during EIS analysis, the non-ideal response of the corrosion system is represented by a constant phase element to obtain accurate impedance values. These are due to geometric distributions, such as surface inhomogeneities and porosity of the electrode.

Figure 7 shows plots from electrochemical impedance spectroscopy of Mg-xLi-1Zn-0.5Ca. The Nyquist plot exhibits different capacitive arcs for the different alloys with relative diameters indicative of their corrosion resistance. The polarization resistance which is clearly seen in the lower frequency region of the impedance bode plot, demonstrates whether an electrode/alloy is reactive or blocking. Furthermore, changes observed in the phase angles illustrate the distinct dielectric behavior of each alloy. In addition to alloy composition, thermomechanical processing plays a major role in degradation kinetics. For example, FSP of dual-phase Mg-Li-Al-Zn alloys resulted in a synergetic strength ductility corrosion optimization by a combinatory contribution of fine grains and nanosized precipitated which inhibited microgalvanic corrosion [26]. A study on optimized Mg-Li-Al-Zr-Y extruded alloys [27] showed decreasing dissolution kinetics as we move from water quenched (WQ) to artificially aged WQ (WQA) to cold rolled WQA (WQAR), as such optimization is required to achieve the best candidate alloys decreasing anodic kinetics compared to pure Mg. The difference in electrochemical response is dependent on the microstructure of the alloy. As the alloy moves from hcp to hcp + bcc to fully bcc, oxide formation, and surface film coverage varies. In the single hcp phase, the surface protection is made up of Mg oxides and hydroxides, which offer poor coverage. However, in the dual phase, there is a development of more stable Li oxide and carbonate, but in the fully bcc phase, complete coverage is achieved with a uniform thick Li carbonate outer layer.

Figure 7.

Electrochemical impedance plots (a) Nyquist plot, (b) impedance bode plot, and (c) phase angle bode plot of Mg-xLi-1Zn-0.5Ca alloys.

3.3 Immersion tests for accessing corrosion behavior

Immersion tests require immersing the alloy in physiological medium for a specific amount of time. This is important because it provides the opportunity to monitor the degradation process and provide direct observation of the alloy at specific times. The hydrodynamic condition of the test setup, among other things, influences the degradation behavior of the alloys. Immersion tests can thus be classified as static, semi-static, or dynamic. For static immersion, degradation occurs in the same media for the specified test duration. On the other hand, dynamic immersion involves the constant flow of the immersion media. Finally, semi-static immersion allows for periodic changes in immersion media.

Semi-static immersion provides a good balance between the static and dynamic test setups, good volume-to-area ratios, and immersion time to avoid distortion of test conditions. In addition, the periodic media changes mimic fluid changes at implantation sites and prevent passivation of Mg alloys at elevated pH values. For semi-static immersion, mass transfer is controlled by migration, diffusion, and a certain degree of convection generated by hydrogen evolution [28].

The determination of corrosion rate can be achieved by volume loss, mass loss, and hydrogen evolution. These methods can generate different results due to testing and measurement limitations such as incomplete removal of corrosion materials during mass loss calculation. The volume loss method can be used both in vivo and in vitro while the hydrogen evolution method can be carried out using a eudiometer system which has to be properly calibrated and well-sealed to avoid the escape of hydrogen atoms. However, mass and volume loss methods have provided consistent results [29]. Chemical cleaning using reagents such as chromic acid is recommended for corrosion product removal to attain uniform weight measurement.

Semi-static immersion on Mg-Li-based alloys reported in Section 2.1 for mass loss after 8 and 90 days is shown in Figure 8. A reduction in degradation rate for L8 and L11 is recorded with increased immersion time, with L8 and L4 having the lowest and highest degradation rates, respectively. Scanning electron microscopy micrographs show the post-immersion morphology of the alloys after 8 days in Figure 9. There is the presence of flakes and multidimensional cracks, and L0 exhibited some islands (Ca₂Mg₆Zn₃ rich region) surrounded by Mg-rich zones. Furthermore, Mg-rich pits demarcated by the Mg₂Ca region were observed. Alloy L4 also displayed pits with high oxide content on the periphery and showed increased pH indicative of an accelerated electrochemical reaction (see Figure 8d). The formation of oxides of MgO, Li₂O, and CaO with PBR <1 is responsible for surface cracking resulting from tensile stresses in the oxide films. This allows the penetration of H₂O and CO_2, leading to the formation of more stable hydroxide and carbonate films. Oxides formed on alloy surfaces can be assessed with tools such as Fourier transform infrared spectroscopy (FTIR), as shown in Figure 8c. The formation of more stable oxides reduced degradation kinetics for Mg-Li alloys in the dual and β-Li phases due to the formation of a passive Li₂CO₃ film.

Figure 8.

Degradation rates of Mg-xLi-1Zn-0.5Ca alloys after immersion for (a) 8 days, (b) 90 days, (c) FTIR spectra of degraded alloys, and (d) pH measurements after 90 days.

Figure 9.

Post-immersion morphology of (a) Mg-1Zn-0.5Ca, (b) Mg-4Li-1Zn-0.5Ca, (c) Mg-8Li-1Zn-0.5Ca, and (d) Mg-11Li-1Zn-0.5Ca alloys after 8 days.

A proper electrochemical assessment of biodegradable implant materials cannot be overemphasized. This is because uncontrolled degradation can lead to pitting, which alters the microstructural composition and mechanical properties of the alloy. Furthermore, when electrochemical redox reaction occurs at an accelerated rate, as is the case with localized degradation, there is a larger amount of ionic dissolution. Such uncontrolled release may result in tissue overload and cytotoxicity.

4. Biocompatibility

Biocompatibility is fundamental for any type of materials intended for use within the human body. When devices are implanted either for diagnosis or therapeutic cases, certain reactions can be generated because of the interaction between the body and the foreign material. These reactions are dependent on the physiochemical properties of the implanted device as well as the implantation site. Generally, the biocompatibility of implant materials is influenced by their surface properties such as surface morphology, chemical composition, surface charge, corrosion rate, thickness, and the nature of the passivation layers.

It is important that metallic implants are biocompatible and non-toxic. In addition, bioresorbable implants also need to degrade in a fashion that does not cause tissue overload arising from very fast degradation. Biocompatibility analysis can be conducted in vivo (as in the case of animal studies) or in vitro using cell lines. It is essential that the implant materials are assessed with cell lines that are related to their intended implantation sites. Materials developed for applications such as cardiovascular use can be assessed with endothelial cell lines, whereas those for orthopedic use can be assessed with osteoblast cell lines.

Different types of assessments can be employed for biocompatibility studies. Commonly used tests include wettability, cytotoxicity, hemocompatibility, and antibacterial responses.

4.1 Wettability

Wettability is a property which influences material-cellular interaction and can be determined by contact angle measurements. It involves measuring the angle a drop of solvent makes with the solid material substrate. Contact angle measurements can be used to determine hydrophilic or hydrophobic properties, surface adhesion, energy, and adsorption via several methods such as sessile or pendant drop. The contact angle differs for various solvents such as water, ethylene glycol, etc., and is dependent on the surface finish and characteristics of the substrate material. Equilibrium contact angle, θ, can be determined from Young’s equation:

$${\gamma }_{lv}cos\theta ={\gamma }_{sv}-{\gamma }_{sl}$$E6

where γ_sv and γ_lv are the surface energy of the solid and liquid, respectively, and γ_sl is solid–liquid interfacial energy.

4.2 Cytotoxicity

A cytotoxicity assessment makes it possible to investigate to what extent a foreign material can coexist with living tissues without causing any toxic effect. Even when bioresorbable implant materials are designed from non-toxic elements, their biocompatibility responses should be assessed as degradation and wear rate can also result in toxicity [30]. A cytotoxicity analysis involves exposing human cells to a test material either directly or indirectly to its dissolved constituent ions to ensure that the implant will not generate high concentrations of ions that can harm surrounding tissues when implanted.

Figure 10 shows cytotoxicity analysis of Mg-xLi-1Zn-0.5Ca alloys on human umbilical vein endothelial cells (HUVEC) with different incubation times. There is increased metabolic activity and cell proliferation with increased incubation time. Other studies on cytotoxicity analysis for Mg-(3.5, 6.5) Li-(0.5, 2, 4) Zn alloys conducted on vascular smooth muscle cells (VSMC) and HUVEC [31] showed that apoptosis was more prevalent in VSMC given the pH and ion concentration of extracts used for the culture. This shows that significant changes in responses can be achieved with different cell lines.

Figure 10.

Cell viability responses of Mg-xLi-1Zn-0.5Ca alloys on HUVEC after (a) 48 hours and (b) 7 days incubation period.

4.3 Hemocompatibility

Hemocompatibility can provide insight into interactions that might trigger activation, secretion, adherence and aggregation of platelets, coagulation, and immunological responses when an implant material encounters blood [32, 33]. It plays a major role in thrombogenicity and is dependent on material’s surface properties. Understanding blood interactions is important because they eventually lead to the formation of a thrombus. Thrombosis can be dangerous when a clot forms within a vessel and cuts off supply, which is a leading clinical complication for stent failure. A hemocompatibility assessment investigates the affinity of blood and its constituents to an implant material, which can be measured using platelet adhesion analysis.

5. Conclusion

The development of functional and safe bioresorbable Mg alloys will be useful in the fabrication of temporary implant devices for both orthopedic and cardiovascular applications.

1. Mg-Li-Zn-Ca alloys, as an example of Mg-Li-based implant material, exhibited enhanced ductility and strength, which are important mechanical properties required not only for a medical device’s in-service application but also good formability that favors ease in manufacturing.

2. The alloys exhibited uniform degradation behavior, which is a very important property for temporary metallic implants such as cardiovascular stents, plates, and screws because it mitigates against early loss of mechanical integrity associated with localized or pitting degradation.

3. Increased cell viability and metabolic activity imply their biocompatibility and nontoxicity, which will mitigate clinical complications and instill confidence in the deployment of magnesium-based prostheses.