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Review

Laser-Based Additive Manufacturing of Magnesium Alloys for Bone Tissue Engineering Applications: From Chemistry to Clinic

by
Mohammad Ghasemian Fard
1,
Fariborz Sharifianjazi
2,
Sanam Sadat Kazemi
3,
Hosein Rostamani
4,5 and
Masoud Soroush Bathaei
5,6,*
1
Department of Mechanical Engineering, Babol Noshirvani University of Technology, Babol, Iran
2
School of Science and Technology, The University of Georgia, Tbilisi, Georgia
3
Biomedical Engineering (BME) Department, Nikan Hospital, Tehran, Iran
4
Department of Biomedical Engineering-Biomaterials, Faculty of Engineering, Islamic Azad University, Mashhad, Iran
5
Department of Materials Engineering, Science and Research Branch, Islamic Azad University, Tehran, Iran
6
Department of Mining and Materials Engineering, McGill University, Montréal, QC, Canada
*
Author to whom correspondence should be addressed.
J. Manuf. Mater. Process. 2022, 6(6), 158; https://doi.org/10.3390/jmmp6060158
Submission received: 1 November 2022 / Revised: 3 December 2022 / Accepted: 6 December 2022 / Published: 10 December 2022
(This article belongs to the Special Issue Frontiers in Digital Manufacturing)
Browse Figures
Versions Notes

Abstract

:
Metallic biomedical implants are made from materials such as stainless steel, titanium, magnesium, and cobalt-based alloys. As a degradable biometal, magnesium (Mg) and its alloys are becoming more popular for applications in bone tissue engineering. Mg-based alloys have been found to be biocompatible, bioabsorbable, and bioactive, allowing them to be used as orthopedic implants with a low Young’s modulus. Computer-aided design can be used to design scaffolds with intricate porous structures based on patient-specific anatomical data. These models can be materialized rapidly and with reasonably acceptable dimensional accuracy by additive manufacturing (AM) techniques. It is known that lasers are the most widely investigated energy source for AM’ed Mg, as they offer some distinct advantages over other forms of energy. Recent studies have focused on developing biodegradable Mg scaffolds by using laser-based AM techniques. In this paper, we aim to review the recent progress of laser-based AM for Mg alloys and survey challenges in the research and future development of AM’ed Mg scaffolds for clinical applications.

1. Introduction

It has been proven that magnesium (Mg) alloys are a viable degradable biomaterial that can be used in orthopedic surgery, respirology, cardiology, and urology [1,2,3]. Mg has the advantage of minimizing or avoiding long-term complications due to the fact that it completely absorbs away over time in the human body [4]. About 65% of the Mg element in the body is found in bones, making it a crucial nutrient for the human body [5]. Additionally, Mg’s modulus is more similar to bone compared to the conventional metallic implants such as stainless steel, CoCrMo alloy, and titanium alloys, which minimizes stress shielding effects in orthopedics [6]. Table 1 compares the mechanical properties of Mg implants with those of typical metallic and polymeric biomaterials [7,8]. Generally, Mg-based alloys are produced by pressure die casting, and wrought Mg-based materials have limited engineering applications, particularly because they cannot be formed and processed at room temperature [9]. Moreover, conventional production methods of Mg alloys for bone tissue engineering applications are challenging, since they cannot generate well-connected pore structures and customized geometric forms, which are the required parameters of novel implants and the personalized medicine field [9,10]. Nowadays, the biomedical industry continues to face many challenges due to the time-consuming and energy-intensive nature of conventional production processes such as casting, forging, and powder metallurgy, especially in the manufacturing of complicated Mg-based implants of high quality [11,12].
The additive manufacturing (AM) of Mg implants is of increasing interest within the biomedical engineering community due to the capabilities of this form of manufacturing, which are not achievable with conventional manufacturing processes, as well as the possibility to design novel biomaterials [13,14]. In addition to the freedom to design (and to optimize topology), additive manufacturing reduces energy consumption and waste by reducing waste materials [15,16,17,18]. It also overcomes the limitations associated with conventional fabricating methods (formative or subtractive) [14]. Through topology optimization and the use of free space as a design variable, the lightest engineering metal can be made even lighter [19]. It is also recommended that components that have a large surface area be used as biomaterials in order to enhance the development of cells, proliferation, and bone regeneration; or as electrodes for Mg in order to provide significant reaction areas [20]. As shown by Equation (3), Mg hydroxide and hydrogen gas are the primary products of the degradation of Mg in water. The following reactions demonstrate the anodic (for Mg) and cathodic (for H2O) nature of this reaction [21]:
Mg → Mg2+ + 2e− Anodic reaction
2H2O + 2e− → H2 + 2(OH−) Cathodic reaction
Mg + 2H2O → Mg(OH)2 + H2 Net reaction
There are a number of challenging problems associated with Mg implants, including their high corrosion rate and low bioactivity. These issues should be addressed before they are used in clinical applications. It is highly desirable to modify the surface of implants containing Mg alloys, so that the corrosion rate is reduced and the bioactivity is enhanced, thus improving their biocompatibility [22].
Using AM technologies, it should be possible to manufacture patient-specific and topologically optimized biodegradable Mg-based implants for bone tissue engineering applications [23]. Up to now, the various AM techniques involved in the fabrication of Mg implants include powder bed fusion, wire arc additive manufacturing, paste extrusion deposition, friction stir additive manufacturing, and jetting technologies. These techniques are shown in Figure 1 schematically [23,24,25,26,27].
Despite the mentioned advantages, there has been limited research in the field of AM’ed Mg implants [25]. The main challenge is that the Mg powder may be subject to oxidation, evaporation, and handling issues as a result of its reactive nature (in atmospheric conditions) [25,26].
Figure 2 shows porous WE43 Mg alloy cubes in three different sizes developed by the authors using the LPBF technique. This work reviews the recent progress on the LPBF of Mg alloys for the construction of bone scaffolds and compares their mechanical corrosion and bio-compatibility. The risk controls employed during laser-based AM techniques have demonstrated excellent success in facilitating the routine and reproducible use of Mg powder additive methods for the manufacture of Mg implants of a wide range of compositions [28,29]. As one of the most powerful and efficient processes in the modern AM field, laser powder-bed fusion (LPBF) has recently been recognized as an efficient and powerful way of fabricating complex 3D objects with a high level of accuracy and reproducibility, as well as superior metallurgical and mechanical properties [29]. Currently, the most commonly used Mg compositions in the LPBF method are pure Mg, Mg-Al, Mg-RE (rare earth), and Mg-Zn alloys, primarily for their comparatively higher market demand, better printability, and properties for use in structural and biomedical applications.

2. AM of Mg for Bone Tissue Engineering

Adult humans are made up of 206 to 213 bones that make up the skeleton as a dynamic support structure [30]. The human body is reinforced structurally through the bone, which allows the body to move, facilitates the process of respiration, allows for adequate breathing, and protects the vital organs within the body [31,32]. Bones also store magnesium, calcium, and phosphorus and produce white and red blood cells as well as platelets [33]. As people age, their bones undergo constant reshaping and biological remodeling [34]. The structure is pulled by involuntary muscles that contract and relax in response to the movement of the involuntary muscles [35]. The structural bones can be divided into three main parts including the macro-, micro-, and nano-scale levels, as shown in Figure 3. Bones can be classified as trabecular or cortical, based on their structure and density [30]. Based on location and age, trabecular and cortical bones have a porosity of 50–90% and 3–5%, respectively [36].
Bone regeneration at the site of damage should be promoted by successful implants that restore the physical function of bone [30]. In general, they should possess characteristics that are similar to those of native bone. These properties include: (i) the mechanical properties should be matched to support loading and reduce/minimize the level of stress-shielding effect; (ii) biocompatibility, suited for cell adhesion/spreading, and does not release toxic ions to extracellular medium; (iii) an extremely porous structure that facilitates bone ingrowth and enables robust biological fixation of the implant to the bone [33]. The natural degradability, biocompatibility, and osteoporotic nature of Mg and its alloys has led to their rapid development as revolutionary metallic biomaterials in the field of bone tissue engineering [3,4,5]. Mg and its alloys have a relatively similar mechanical strength to that of human bones, as seen in Table 1. Mg and it alloys have high biocompatibility and similar mechanical properties with natural bone properties that meet two requirements of implant design intrinsically [7,8,9,10]. Moreover, the permeability of the porous metal used in bone implants is also an important characteristic as a third requirement, since blood must travel through them in order to transport cells [36]. By utilizing the AM technique as a new technology, porous Mg-based scaffolds that are tailored to the anatomical needs of a patient can be manufactured layer-wise using well-developed architectures derived from computer-aided design (CAD) models in accordance with the patient’s anatomical requirements [38]. Unique abilities that could not be reached in any way in traditional Mg production methods can be used, such as die cast [39]. It is also possible in AM to adjust the modulus of the implant or porous mesh by varying the porosity and strut size [40]. Among the types of patients who would benefit from AM strategies are those who require a resection of bone tumors, treatment following fractures, and other types of bone loss, whether regular or irregular. It has been established that LPBF, formerly known as selective laser melting (SLM), is the most widely explored method for the AM of Mg and its alloys [13,14,25], with very few attempts to use direct laser deposition (DLD). Therefore, DLD was excluded from the main body of the present review.

3. Laser Powder-Bed Fusion (LPBF)

Using CAD design, laser processing, and computer numerical control, PBF creates metallic components with enhanced properties [41,42,43]. Figure 4 illustrates the typical LPBF manufacturing process: (i) the computer obtains the model slice data from the 3D model by converting it to an STL file; (ii) the forming cylinder is covered with a layer of Mg powder that is evenly distributed; (iii) as a result of the computer’s control of the laser beam based on the model slice data, the powder layer is imaged selectively with the laser beam, which receives heat and starts to melt/solidify, resulting in a single layer of material; (iv) a new layer of powder is paved by the roller after the molding cylinder is lowered by a layer. After the first layer has been constructed, the second layer will be applied on top of it, until the final part has been formed [44,45]. In LPBF, Mg powders varying from 15 μm to 160 μm are used, with the majority of powder being on the lower end of the mentioned range [46,47]. Moreover, the chemical composition, mechanical properties, and geometry of manufactured Mg implants can vary due to LPBF’s wide range of parameters. Laser power, scanning speed, and layer thickness are the most important parameters in the LPBF process [48,49,50].
Most of the experiments on the PBF of Mg alloys have been conducted using lasers [50,51,52]. The lasers concentrate high concentrations of heat over a limited area of the powder bed, melting the powder over a limited timeframe [53]. Then, rapid solidification occurs as a result of this short-timed heat flux. It is possible to carry heavier loads due to this rapid solidification, which refines the grain compared to conventional production methods [54,55]. The vaporization of some elements of Mg-based powder occurs when the powder is heated to a high temperature [56]. In the melt pool, powder vaporization builds up vapor pressure locally [57]. In a melt pool, molten material is spattered outward by pressure, forming low-density structures that may reduce the stress-shielding effect [57,58]. In addition to variations in the chemical composition, this also leads to variations in particle size [59]. For printed components to perform well under corrosion conditions, alloying elements must have good solubility during AM and create corrosion resistance intermetallic phases [59]. PBF is influenced in a significant way by the power of the laser and scanning speed in order to determine the melt pool, vaporization, and the deposition results [60]. There are many factors that affect the quality of depositions, including laser power and scanning speed [61]. However, the effects of each factor on the other are difficult to describe on their own [60]. In combination, they are essential for determining how much energy is transferred to Mg powder via the laser [60]. It has been reported that spherical Mg powder particles with a purity of 99.9% and a mean size of 24 μm were used with 155.56 J/mm3 energy density to achieve 97.5% density [62]. Moreover, researchers found that in response to an increase or decrease in energy density, the material’s relative density and mechanical strength deteriorated [63]. In the case of Mg alloy powders, it was found that Mg and zinc elements within ZK60 powder were heavily vaporized when subjected to 1250 J/mm3 of energy density [56]. The melt pool stabilized and the vapor pressure was reduced after the laser density was reduced to 250 J/mm3 [64]. A poor relative density of 82.25% was the result of the incomplete fusion of powder particles. With an energy density of 416.67 J/mm3, a maximum relative density of 94.05% could be achieved, as seen in Figure 5a [65]. In another case, with an energy density of 238 J/mm3, WE43 was printed at a relative density of 99.4%. WE43 is an alloy of Mg that contains the primary alloying elements yttrium and neodymium [66]. There is evidence suggesting that optimal printing parameters are generally found when the energy density of the melt pool is low, resulting in a high part density and low vaporization of the alloying elements inside the melt pool [67]. In addition, porosity increases with increasing energy density. The porosity increased from 0.4% to 17% when the laser power was decreased from 195 W to 135W, and the scanning speed increased from 800 mm/s to 1200 mm/s [68]. In split Hopkinson pressure bar tests, the dynamic strength decreased with decreasing energy density [69]. A porous scaffold composed of diamond unit cells could also be constructed with a WE43 alloy. Unit cells with a low energy density of 100 J/mm3 were able to achieve strut sizes of 400 μm [60]. In another study, the laser power was 15 W, and the scanning speed was 20 mm/s, which allowed the Mg-Al alloy to melt completely (Figure 5b) [70]. This is equivalent to 187.5 J/mm3 of laser energy. The Mg–9%Al alloy may have a range of acceptable energy density regions based on a study of the same material at an energy density of 155.6 J/mm3. In addition to powder quality and layer thickness, several parameters affect this range of energy density [71].
LPBF of metallic materials must also take porosity into account, which is a common problem [72]. An analysis of how processing parameters affect the porosity of Mg specifically has been presented in the following Table 2. Changing any one of the listed parameters will result in a different porosity, so it is difficult to extract a trend based on the effect of any individual processing parameter [72]. In order to achieve high density and low porosity, alloy composition dictates the ‘printability’ of the alloy during the optimum processing windows [73]. It is inevitable that porous materials will tear and crack; however, it is necessary to avoid hot tearing or cracking while porous materials are present [74]. A LPBF component’s quality is deteriorated most severely by hot tearing and cracking. When the temperature gradient remains high, even when the constitutional super-cooling is lower, columnar grains are more susceptible to hot tearing [75]. Cavities and hot-tearing cracks can occur across the full length of columnar grains when temperature and liquid volume fraction decrease. Since there has been limited research on the impacts of alloying and processing parameters on hot tearing during LPBF in Mg, it is unclear what effect they have [76].

4. The Advantages of LPBF over Other AM Processes

It is worth noting that LPBF has the added advantage of being able to produce highly porous and fine structures as well as accommodate many forms, not only prismatic ones [72]. Thus, it is widely used to manufacture metallic scaffolds and implants. In addition, osseointegration studies have shown that implants can be designed with preferentially porous layers to enhance adhesion between the bone tissues and implant material [75]. A human bone consists of three distinct anatomical cavities: haversian canals, osteocytic lacunae, and canaliculi [66,67,68,69]. A bone’s mechanical characteristics and processes are remodeled in these three cavities. Living tissues can flourish in porous structures, as nutrients can be transported [49,50,51,52,53,54]. It is possible to design components with a modulus similar to the bone’s, which reduces stress-shielding problems in metallic orthopedic materials by adjusting porosity levels. Based on the geometries of metallic cellular structures, non-stochastic and stochastic structures are distinguished [77]. An arbitrary number of closed or open voids characterize stochastic metal porous structures. The non-stochastic structure of metal shows homogeneous patterns that repeat over many units cells, as opposed to the stochastic structure [78]. Because metal cellular/porous structures are lightweight and have better power absorption properties and excellent acoustic and thermal properties, they are commonly used in the automotive, aerospace, and chemical industries [79]. A structure with such complex external shapes and intricate internal structures is difficult to fabricate using traditional casting and powder metallurgy techniques. Changing the specifications of these production techniques may alter the dimensions of the pores, but it will only produce a porous structure that is arbitrarily arranged [80]. Using LPBF, on the other hand, can create porous metals with a predefined external form and an internal structure that complements the rigidity of bone and reduces or diminishes the need for stress shields during manufacturing [81]. Compared with directed energy deposition (DED), which is also widely used for the fabrication of metals, PBF has a smaller beam spot, finer powder, and a thinner layer, which provide better dimensional accuracy [10]. In addition, the LPBF process is characterized by its high energy density and the absence of sacrificial binders, which allows for nearly complete densification of metal parts, which is an advantage over other metal additive manufacturing technologies such as binder jetting. Laser and electron beam spot sizes are currently between 50 and 100 microns, powder sizes are between 20 and 50 microns, and the thickness of each powder layer is between 20 and 80 microns for LPBF and electron beam powder bed fusion (EBPBF), respectively. Therefore, the dimensional deviation between the as-built geometry and the designed geometry is decreased as a result of the higher forming accuracy of LPBF when compared to EBPBF [13]. Its high dimensional accuracy, high performance, geometric freedom, and geometric freedom without rigid support make LPBF suitable for metal implants. There are many disadvantages to the wire arc additive manufacturing (WAAM) of Mg in comparison with LBPF, including the possibility of scaffold structures that might be too coarse to be used in medical applications, the availability of wire (no custom alloys), and the high level of processing required. Moreover, compared to LPBF, WAAM produces components with a high deposition rate and surface roughness, resulting in a higher degradation rate of WAAM’ed Mg implants in physiochemical solutions [14].

5. Properties of LPBF’ed Mg Implants

5.1. Mechanical Properties

Mg components processed using LPBF have a significantly improved mechanical performance due to the thin continuous refined microstructure formed [10]. The ideal powder particle size should be between small and large: finer particles can be melted easily and are beneficial to model quality surface finishes, and larger particles are beneficial for ductility, hardness, mechanical strength, and toughness [13]. Further, components produced by LPBF typically exhibit anisotropic microstructures at a variety of length scales. A rapid solidification process forms an anisotropic microstructure in the heat dissipation direction through conduction, convection, and radiation. It is important to consider the type of scanning strategy used, the base plate temperature, and the direction in which the build takes place when determining anisotropy. A component built in a different direction to the substrate such as perpendicular (e.g., vertical) or parallel (e.g., horizontal) has a different thermal history [13,14]. This creates an anisotropic mechanical property as well as a different surface texture (finish). Selecting the right combination of parameters is crucial to the quality of LPBF’ed Mg implants. As a result of non-optimized LPBF process parameters such as laser energy density, metallurgical defects, cracks, and pores are formed, resulting in poor mechanical properties. The tensile properties of LPBF’ed Mg and its alloys are listed in Table 3. The yield strength of LPBF’ed Mg alloys compared with cast and wrought (rolled and extruded) alloys is shown in Figure 6. Typical alloy yield strength for most structural applications is greater than 200 MPa, with some alloys reaching 350 MPa. A majority of LPBF’ed Mg alloys have less than 5% ductility, and some alloys may not even have ductility. Compression or hardness tests are required to measure their mechanical properties. The low ductility of such a material makes it unsuitable as an engineering material. Several alloys are characterized as having fine grains, weak textures, and low porosities, yet ductility is still poor [87]. Here are some reasons for the change. Because the as-LPBF solidifies rapidly, it has a high residual stress, which reduces its ductility. There is a high concentration of intermetallic elements along the grain boundaries of alloys such as the AZ91, WE43, and Mg-Gd alloys studied [70]. Thereafter, the grain boundary becomes brittle and responsible for local failures (e.g., slip and twin transmission across grain boundaries, grain boundary sliding, etc.). Once this happens, they do not have the ability to accommodate plastic deformation. It would also be possible for the sputtering powder or vapor to re-deposit on the surface of the sample, which would result in the part being poorly consolidated or with a weak bond to the sample [13].

5.2. Corrosion Behavior

The corrosion of metallic implants develops rapidly in the human body as a result of aggressive conditions. As a result of the chloride ions in biological media, corrosion occurs, and H2 gas is produced, which slows down the healing process and causes necrosis of the tissues in the surgical area [97]. Surface coatings can be applied to the surface of Mg implants to inhibit the corrosion progress [98]. The LPBF’ed Mg alloys are currently used most frequently in biodegradable scaffolds and implants. In the case of oral and maxillofacial implants, for example, the Mg-based biomaterials should remain mechanically intact for the initial month, before gradually decomposing and eventually dissolving. Due to the poor corrosion resistance of Mg and its alloys in most aqueous environments, enough electrochemical durability is needed. Table 2 illustrates the electrochemical corrosion behavior of LPBF’ed Mg and its alloys in simulated biological solutions. It has been reported that the corrosion current density (icorr) of LPBF’ed pure Mg ranges from 74–177 µm/cm2 in Hank’s solution, which is higher than the casted pure Mg ingot measured under the same condition (23.6 µm/cm2). Through the manipulation of print process parameters, it is possible to alter the porosity of a 3D construct, which has a significant impact on corrosion rates and the behavior of the cells. According to the processing parameters, mass losses range from 3 mm to 32 mm per year. Pure Mg manufactured by LPBF has an extremely high corrosion rate, about 144 mm/year when it is dissolved in a 3 wt.% NaCl solution. Due to the loosely fused Mg clusters and sintered Mg powder, the corrosion resistance is significantly reduced. This higher corrosion rate is closely related to the defects generated during LPBF. Indeed, degradation rates are faster when there are more defects. Furthermore, the corrosion resistance of the LPBF’ed WE43 alloy was much lower than the casted alloy. In comparison with the interface and cast samples, the LPBF’ed WE43 exhibits a significantly higher corrosion potential and significantly lower corrosion current density. Accordingly, all samples exhibit localized corrosion morphologies with differing severity, based on the observed active-like behavior. As a result of the absence of a distinct Tafel region in the anodic polarization curves, the quantitative values deducted from the polarization plots cannot be used individually, but can only be used in qualitative comparative studies of the polarization response [99]. Based on 5% foetal bovine serum in a revised simulated body fluid (r-SBF), corrosion current density ranges from 20 to 60 m/cm2. Moreover, it loses mass 6 to 7.2 mm per year when it is dissolved in 0.1 M NaCl solution, which is approximately six times greater than the mass loss rate of cast WE43 alloy (0.8 to 1.2 mm per year). A high density of stable RE oxide particles and a reactive Mg matrix results in a micro-galvanic reaction, resulting in a higher corrosion rate. If plasma electrolytic oxidation is not performed on the scaffold surface of the LPBF’ed WE43 scaffold after 21 days of immersion in SBF, its structural integrity is reported to be lost after 21 days of immersion in SBF. Research in another study reported that the LPBF’ed WE43 alloy retained structural integrity after 28 days in vitro without obvious particle separation; however, it had a 41% loss in strength. Structural parameters, such as pore size and strut diameter, seem to affect scaffold degradation more than bulk samples [95]. The corrosion resistance of the LPBF’ed Mg-Al alloy is similar to the cast alloy, however. In the SBF state, the degradation rate for LPBF-AZ61 alloy was around 1.2–2.7 mm/year as-immersed, and then it decreased. After 24 days of immersion in SBF, the corrosion rate of cast AZ61 samples stabilizes at 1.299 mm/year, which is similar to the corrosion rate for the cast AZ61 sample. Although the surface of the corroded sample seems to indicate a more severe corrosion of LPBF’ed ZK60 than cast ZK60, the LPBF’ed ZK60 has marginally better corrosion resistance in Mg-Zn systems than cast ZK60. By blending ZK powders with Cu powders, Shuai et al. increased the antibacterial effect of Mg-Zn-Zr implants by adding dilute Cu concentrations to LPBF ZK30 and ZK60 [100]. LPBF’ed ZK-Cu alloy degradation is accelerated by Cu addition.

5.3. Biocompatibility

As AM’ed Mg alloys are the most promising materials for bone tissue implants, there is a need to consider LPBF’ed Mg alloys’ biocompatibility. The human body requires Mg to function properly. Due to its biodegradability, Mg is capable of gradually transferring a load from the implant to the regenerated bone if the implant is biodegradable enough [101]. Mg has a degree of biocompatibility and bioactivity that allows it to promote the proliferation and differentiation of cells even though it is biocompatible [102,103]. In addition to stabilizing DNA and RNA, it promotes bone growth and healing. Biocompatible alloying elements are also needed when designing bioabsorbable biomaterials based on Mg alloys [104]. However, a rapid breakdown of Mg scaffolds provides significant amounts of hydroxides, Mg ions, and ions from alloying elements to be released in large quantities, which have an adverse effect on the viability of cells and biocompatibility of the alloy. The study of cell interactions with biomaterials can be performed by culturing cells [105,106,107]. LPBF’ed WE43 has been found to be biocompatible in vitro in numerous studies [108]. Four hours after immediate seeding, MG-63 cells in direct contact with WE43 appeared to be dying as determined by live-dead staining, followed by dual-channel fluorescent optical imaging (FOI) [109]. After seeding Ti-6Al-4V, on the other hand, no cell death can be observed after 4 h. It was found that a substantial number of cells survived 24 h after direct contact with WE43 scaffolds that had been pre-incubated in physiological serum-containing culture medium for 48 h [110]. Using a similar methodology, it was reported that poor cellular adherence to the scaffold was observed in a similarly designed study using indirect extract-based assays (LDH, XTT, and BrdU). RE-based Mg alloys themselves appear to possess no cytotoxic potential; however, the vast reaction on their bare metal surfaces causes high levels of hydrogen gas to be evolved and pH shifts local to the surfaces, impairing the metabolic efficiency of the cells [111]. LPBF’ed WE43 scaffold was found to contain no viable cells, and only a few dead cells were visible by direct live/dead staining. There is a possibility of resolving this problem if the surface of the substrate can be modified, using plasma electrolytic oxidation, which seems to be a suitable niche for adhering cells when the surface is passivated [112]. Additionally, this technique appears to have reduced degradation byproduct releases, thus indicating that hardly any signs of cell impairment were observed. Additionally, there are reports that LPBF scaffolds are manufactured on pre-alloyed Mg-Nd-Zn-Zr (JDBM) materials, as shown in Figure 7a–i [113]. Similarly, in the trial of cell adhesion to scaffolds coated with dicalcium phosphate dihydrate vs. scaffolds left uncoated, more cells adhered to the scaffold coated with dicalcium phosphate dihydrate. Neither the coated nor uncoated LPBF’ed JDBM scaffolds showed a significant difference in terms of cytotoxicity in this study: both examples stimulated cell proliferation (Figure 7j–o). The findings presented in this study and the verification of these findings require a fuller study and verification, because it is highly unlikely that a direct cell response will not at least be induced by uncoated AM’ed Mg scaffolds.
Moreover, the presence of bacteria in the implant may cause the implant to loosen or even fail, demonstrating the importance of inhibiting infections caused by bacteria [114,115,116,117,118]. The positively charged Cu ions released by the body are capable of attaching to the negatively charged bacterial cell walls, impairing the permeability of the membranes and eventually causing the bacterial cell to lyse and die. Additionally, Cu ions can also activate some enzymes by interacting with their thiol groups, resulting in bacterial death. Moreover, copper functions as a cofactor in a wide range of metabolic enzymes in the human body [119]. LPBF’ed ZK60-Cu alloys prepared by Xu et al. exhibited strong antibacterial properties when they contained Cu [95]. This study indicated that ionized Cu ions, combined with an alkaline environment, could effectively kill bacteria by damaging the structure of cellular membranes, causing enzymes to denaturate, inhibiting deoxyribonucleic acid replication and destroying DNA.

6. Conclusions, Challenges, and Future Perspectives

The biocompatibility and biodegradability of Mg alloys make them ideal materials for biomedical applications. Due to the manufacturing feasibility of complex-shaped components, it would be possible to fabricate customized Mg components via additive manufacturing techniques based on patient-specific anatomical data. The present review article discusses the PBF-based AM process, which is used to fabricate Mg-based implants. The reasons behind its selection over other AM techniques are compared and discussed in detail. The following is a concise summary of the challenges and opportunities that emerge in light of the review presented herein.

6.1. Bottlenecks

(1)
Laser-based additive manufacturing presents a challenge in terms of producing pre-alloyed powder. More research is needed in the area of blending magnesium powders and building consistency.
(2)
How the topology of scaffolds affects cell proliferation, new cell growth, and the lattice structure of the fabricated Mg components that are fabricated using the additive manufacturing process are still unexplored. This suggests that further studies should be conducted using in vitro and in vivo methods for the Mg scaffolds manufactured through LPBF.
(3)
Mg implants are evaluated in vivo for their biodegradation performance out of both processes. Therefore, the study of in vivo processes should be carried out in great detail in order to succeed in clinical applications. For biomedical implants, LPBF of Mg components has been shown to be an appropriate and promising alternative. An alloy that is suitable for bio-implant application could be developed by evaluating the Mg alloys used currently. An alloying element would be added according to the strength considerations of the implant in question and its biocompatibility. This will be considered in future work if a new Mg-based alloy is created.

6.2. Prospects

(1)
As a result of the efficient infiltration and complete melting of Mg alloy, LPBF is a suitable AM technology for the fabrication of Mg implants. This resulted in the removal of voids and the creation of high-density components. The powder properties that were used in the manufacturing process of the Mg scaffolds and implants, as well as the printing parameters used in printing, play a major role in determining their biological and mechanical properties.
(2)
LPBF produces the Mg scaffold with a hierarchical porous structure that mimics the structure of the human bone in terms of micro- and macro-pores for personalized medicine.
(3)
In comparison with other AM techniques, LPBF provides better dimensional accuracy, because it has a smaller beam spot, finer powder, and a thinner layer. Additionally, LPBF technology offers high-energy density, no sacrificial binder, and near-complete densification of metal parts, which makes it superior to other metal additive manufacturing processes such as binder jetting and WAAM.
In conclusion, it is expected that future developments of the LPBF process of Mg, particularly those for bone tissue engineering applications, will focus on a wide range of properties, including mechanical properties, corrosion performance, biocompatibility, printing properties, biomimetic properties, and biodegradation properties. It would be a reliable method for building various organs and tissues with diverse mechanical requirements whether the use of new types of steels as biomaterials with different compositions is being developed. It would be a reliable way to achieve reprogrammed mechanical properties and functions. The Mg-based alloys produced by LPBF will be used to make stents, screws, plates, and scaffolds in the near future.

Author Contributions

M.G.F.—conceptualization, methodology, and writing (original draft preparation); F.S.—resources and writing (original draft preparation), supervision; S.S.K.—software, writing (original draft preparation); H.R. writing (original draft preparation); M.S.B.—writing (review and editing) and supervision. All authors have read and agreed to the published version of the manuscript.

Funding

This research received no external funding.

Data Availability Statement

Not applicable.

Conflicts of Interest

The authors declare no conflict of interest.

References

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Jmmp 06 00158 g001 550
Figure 1. Schematic of (a) powder bed fusion [23], (b) wire arc additive manufacturing [24], (c) paste extrusion deposition [25], (d) friction stir additive manufacturing [26], and (e) jetting technologies [27].
Figure 1. Schematic of (a) powder bed fusion [23], (b) wire arc additive manufacturing [24], (c) paste extrusion deposition [25], (d) friction stir additive manufacturing [26], and (e) jetting technologies [27].
Jmmp 06 00158 g001
Jmmp 06 00158 g002 550
Figure 2. The WE43 Mg alloy cubes in three different sizes developed by Dr. M. S. Bathaei’s group using the LPBF technique for bone tissue engineering applications.
Figure 2. The WE43 Mg alloy cubes in three different sizes developed by Dr. M. S. Bathaei’s group using the LPBF technique for bone tissue engineering applications.
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Jmmp 06 00158 g003 550
Figure 3. Hierarchical macro- to nanostructures of natural human bone [37].
Figure 3. Hierarchical macro- to nanostructures of natural human bone [37].
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Figure 4. Mg-based bone implant manufacture using a typical LPBF technique that uses laser-based AM principle [44].
Figure 4. Mg-based bone implant manufacture using a typical LPBF technique that uses laser-based AM principle [44].
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Jmmp 06 00158 g005 550
Figure 5. (a) Influence of laser scanning speed on relative density of ZK60 [65] and (b) grain size variation of Mg–9%Al powder as a function of laser power and scan speed [70].
Figure 5. (a) Influence of laser scanning speed on relative density of ZK60 [65] and (b) grain size variation of Mg–9%Al powder as a function of laser power and scan speed [70].
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Jmmp 06 00158 g006 550
Figure 6. Comparison of tensile properties of Mg alloys developed by LPBF and other conventional production techniques [87].
Figure 6. Comparison of tensile properties of Mg alloys developed by LPBF and other conventional production techniques [87].
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Figure 7. (a–c) Micro-CT images, (d–f) SEM images, and (g–i) corresponding magnified images of the as-polished B (a,d,g), D (b,e,h), and G (c,f,i) scaffolds. Cytocompatibility of (j–l) the G scaffolds and (m–o) the G-DCPD scaffolds, where (j,m) show the results after 6 h, (k,n) show the results after 1 d, and (l,o) show the results after 3 d [113].
Figure 7. (a–c) Micro-CT images, (d–f) SEM images, and (g–i) corresponding magnified images of the as-polished B (a,d,g), D (b,e,h), and G (c,f,i) scaffolds. Cytocompatibility of (j–l) the G scaffolds and (m–o) the G-DCPD scaffolds, where (j,m) show the results after 6 h, (k,n) show the results after 1 d, and (l,o) show the results after 3 d [113].
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Table
Table 1. Comparison of Mg implant properties with other biomaterials [7,8,9,10].
Table 1. Comparison of Mg implant properties with other biomaterials [7,8,9,10].
Tissue/MaterialYoung’s Modulus (GPa)Yield Strength (MPa)Compression Strength (MPa)Tensile Strength (MPa)
Cortical bone7–30-100–230164–240
Cancellous bone0.01–3.0-2–12-
Ti-6Al-4V (casted)114760–880-895–930
Ti-6Al-4V (wrought)114827–1103896–1172860–965
Stainless steel 316 L193170–310480–620540–1000
CoCrMo Alloy240500–1500-900–1540
Mg (99.9%, casted)41214087
Mg (99.9%, wrought)41100100–140180
Table
Table 2. Processing-relative densities of LPBF’ed Mg and its alloys for bone tissue engineering applications.
Table 2. Processing-relative densities of LPBF’ed Mg and its alloys for bone tissue engineering applications.
AlloysPowder Size and Shape (μm)ParametersInput Energy Density (J/mm3)Relative Density
(%)		Ref.
Power (W)Spot Size
(μm)		Speed
(mm/s)		Thickness
(μm)		Hatch Spacing (μm)
MgPre-alloyed 43, s 1901001003010030096.1[77]
90100>300Evaporated
MgPre-alloyed 24, s7080500303015697.5[78]
1250 6388.2[79]
WE4325–63, s195100800302004199.7[67]
1958002503398.3
19512002002796.6
13512002001987.6
WE4325–63, s200125700304023899.9[28]
WE4325–63, s200701100401303599.7[80]
WE4325–63, s20090700304023899.8[81]
WE4330, p 45, 63, s12090960304010498.6[82]
1501200 10499.0
3001200 20899.5
GZ151K25–65, s200–700307013697.9[83]
GZ112K31–44, s801001003010026798.7[65]
3008999.9
5005399.7
7003899.8
10002796.9
15001871.8
5005010799.5
5001503696.5
G10K63, s80–2003010013399.2[84]
Mg-1ZnBlended Mg-5.5 Zn (36, s), Mg (31, s) and Zn (19, s)180150700207018399.4[47]
Mg-2Zn98.2
Mg-6Zn94.7
Mg-12Zn98.9
ZK6030, s501506.710010075094.5[85]
8.360097.4
1050088.6
11.742072.8
ZK6030, s200150300208041794[86]
50025093
70017988
90013984
1 Spherical shape.
Table
Table 3. Input energy density, grain size, tensile, and electrochemical properties of LPBF’ed Mg and its alloys.
Table 3. Input energy density, grain size, tensile, and electrochemical properties of LPBF’ed Mg and its alloys.
AlloysEnergy Density (J/mm3)Grain Size
(μm)		Mechanical PropertiesElectrochemical PropertiesRef.
H 1
(HV)		YS 2
(MPa)		UTS 3
(MPa)		EL 4
(%)		Solutionicorr
(μA/cm2)		Mass Loss
(mm/year)
Mg97.51–5––––Hank’s solution743[77]
88.2–17732
Mg300–52.4–––––[70]
Mg-9Al25010–2070–––––[88]
Mg-9Al1561–3 –2741––[89]
AZ611391.6–2192733.3––[90]
1561.82332873.1
1792.12252612.8
2082.52162392.1
AZ611204.570–––SBF solution–2.7[91]
1408802.4
160109312
18013901.5
AZ911671–2.985–1002742961.2––[92]
832372541.8––
AZ9168.61–10115–––––[47]
AZ911041–1.5–2653283.8––[93]
AZ912783.3–3083451––[79]
AZ91-SiC2781.1–2603002––[67]
AZ91–2Ca–––2353323.2––[29]
WE4312034----0.1 M NaCl5.16–7.2[28]
150275.0
300184.4
WE432381–29630812.2––[80]
WE43351–3–2142512.6––[82]
WE4323820.4––––––[82]
G10K13327801802282.2––[83]
GZ151K1362 3453683––[94]
Mg-1Zn183–5014511––[95]
Mg-2Zn46702.5
Mg-6Zn65501.5
Mg-12Zn83753.2
ZK302000–80–––SBF solution17.81.23[96]
ZK30-Cu9847.82.12
1 Hardness; 2 yield strength; 3 ultimate tensile strength; 4 elongation.
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MDPI and ACS Style

Fard, M.G.; Sharifianjazi, F.; Kazemi, S.S.; Rostamani, H.; Bathaei, M.S. Laser-Based Additive Manufacturing of Magnesium Alloys for Bone Tissue Engineering Applications: From Chemistry to Clinic. J. Manuf. Mater. Process. 2022, 6, 158. https://doi.org/10.3390/jmmp6060158

AMA Style

Fard MG, Sharifianjazi F, Kazemi SS, Rostamani H, Bathaei MS. Laser-Based Additive Manufacturing of Magnesium Alloys for Bone Tissue Engineering Applications: From Chemistry to Clinic. Journal of Manufacturing and Materials Processing. 2022; 6(6):158. https://doi.org/10.3390/jmmp6060158

Chicago/Turabian Style

Fard, Mohammad Ghasemian, Fariborz Sharifianjazi, Sanam Sadat Kazemi, Hosein Rostamani, and Masoud Soroush Bathaei. 2022. "Laser-Based Additive Manufacturing of Magnesium Alloys for Bone Tissue Engineering Applications: From Chemistry to Clinic" Journal of Manufacturing and Materials Processing 6, no. 6: 158. https://doi.org/10.3390/jmmp6060158

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