Medical material development has been prominent over the past three decades. Throughout this time, the most notable materials used were ceramics due to their similar constituents to the human body, for example, the bone-regeneration constituent containing calcium phosphate that can be easily reproduced (Best et al., 2008). Ceramic material is categorized as bioinert or bioactive and shows strong interfacial bonding to the host tissue and osseointegration capability compared to other material classification. When it comes to medical device classification, it is subdivided into three categories implemented by the Food and Drug Administration (FDA). These are Class I, II, and III, with the latter being described as high risk to the human body, as illustrated in Figure 1; therefore, safety measures are adopted and required to be approved prior to mass release to the public (Johnson, 2016).

In the past, there were no systematic guidelines being introduced for materials, such as metallic, ceramic, and polymer, that could be considered for commercial biomedical devices. This resulted in more than 700 deaths and 10,000 patient injuries (Dhruva and Redberg, 2013). As the framework became gradually implemented, likewise for 3D-printed devices, similar validation and specific requirements such as design and quality control strategies came to be utilized, resulting in controlled output and consistent production of the devices (Morrison et al., 2015). Within the European Union (EU), custom-made implants, which are intended to penetrate beneath the epithelial layer such as contact lenses or devices embedded inside the body, require CE marking and have to abide by regulation (Aimar et al., 2019).

To fabricate parts of complex and intricate structures to fit nicely on the patient, for example, skull reconstruction, traditional ways involving mesh implant insertion limit dimensional precision and structural integrity. Therefore, 3D printing has an upper hand when it comes to designing complex structures which allow successful integration to the host tissue. Particularly, in 3D-printed dental implants, the flexibility of tuning micropore channel architecture consequently facilitates signaling molecules to be activated at specific sites. This is useful in recruiting cells and maintaining alveolar height in a dental implant for better healing ability (Yin et al., 2021). The absence of surface treatment in the earlier period of pacemaker implant invention that revolves around metallic material, such as stainless steel, caused challenging mechanical fracture to its “pacemaker lead,” eventually causing declining electrical stimulus. Many attempts have been made to replace the metallic alloy component, but no sign of improvement has been observed aside from more issues, including those related to corrosion (Joung, 2013). There are many significant benefits surrounding 3D printing as mentioned in many articles from its customization (Attarilar et al., 2020) to being cost effective (Schubert et al., 2014), and being ready-built in several hours (Mertz, 2013). Specifically, healthcare professionals use 3D printing to engage with patients via software, such as MIMICS or MeDraw, without requiring engineering expertise to develop and analyze models (Dai and Xu, 2021). Through 3D printing, anatomical models serve to minimize design error on patients and provide extensive practice to physicians prior to surgical procedures. As a result, this provides direct communication between physicians and patients to translate CT/MRI scans beyond a two-dimensional layout when treating life-threatening disease. This multidimensional function of 3D printing thus tackles medical issues effectively and efficiently while reducing the long waiting lists for treatment (Aimar et al., 2019). Another point that is not often emphasized or is overlooked is the open-source nature of the 3D printing files which allows vast collaboration among researchers and physicians more than just restrictive parameters published in scientific journals. The open-source database allows professionals to selectively tailor design dimensions according to the anatomical model of patients (Gross et al., 2014; Ventola, 2014). To a greater degree, open-source sharing of 3D models in the STL (Standard Tessellation Language or STereoLithography) file format has been proactively used in the immunology field, such as in the tiny replicas of microorganisms to visualize proteins and viruses (Coakley et al., 2014). In short, the importance of 3D printing has extended across vast medical applications from operation planning, as instrument guides to implant devices, and even in the field of microorganisms (Dai and Xu, 2021).

A typical procedure of a bottom–up 3D printing approach has been illustrated by Wang et al. (2017) (Figure 2). Particularly in orthopedic application, there are drawbacks in manufacturing porous implants using conventional methods which subsequently lead to implant failure due to lack of bone–implant integration. Depending on the final medical device customization and mechanical profile desired, a variety of surface modification techniques can be adopted, which are explored more in the Surface Modification Techniques section.

Metallic bio-devices can be divided into degradable and nondegradable types. A degradable metallic material, such as magnesium alloy, has bioactive capability and interfacial interaction between implant and tissue. The ability to degrade is preferred for temporary implant support such as plates and screws in bone healing, thereupon eliminating the unwanted surgical risk and excessive cost of a second surgery (Gu et al., 2014). By contrast, a nondegradable metallic material is classified as bioinert through the formation of a tissue fibrous capsule which isolates the implant from the surrounding tissues; however, the process does not trigger adverse interference to the biological system to a certain extent (Daghighi et al., 2013). A suitable implant can be either bioinert or bioactive as long as it does not induce toxicity. In principle, a material serving as an implant in the human body must ensure biocompatibility to the host tissue. In the medical field, the commonly used materials for orthopedic application are comprised of titanium alloy (Xiu et al., 2016), cobalt chromium alloy, and stainless steel (Balazic and Kopac, 2007), which are also categorized as nondegradable materials.

The strength-to-weight ratio of titanium alloy and magnesium alloy demonstrates similar mechanical capability to the bone, consequently exhibiting good biological integration. However, when it comes to clinical capability, these alloys tend to suffer from accelerated corrosion rate and poor cell viability when in contact with human body fluids (Kazantseva, 2018; Zhang et al., 2019; Atrens et al., 2020).

Metals are known to corrode easily but some are more resistant. This is due to the formation of a passivation layer that hinders corrosion from taking place. In general, most metals have low corrosion resistance when compared to noble metals such as gold and silver. When in contact with tissue, due to the oxygen diffusion limit contributed by the fast leakage of metal ions in the body, these ions create a high toxicity level, thereby causing adverse effects to the cells (Steinemann, 1998), as shown in Figure 3A. Corrosion is a common phenomenon and can be avoided through surface oxide modification. A study was conducted to identify cell proliferation between bare titanium alloy and cobalt chromium alloy against hydroxyapatite-coated metallic alloys. It was shown that there was a significant increase in cell viability when being coated (Yuan et al., 2018).

When an implant to the body is inevitable to save one's life, one needs to seriously examine the possibility of an adverse reaction of the implant to the body. Figure 4 shows the tissue responsive effect on various common metallic implant materials. It was evident that metallic implants such as cobalt (Co), titanium (Ti) alloys, and iron (Fe) show a relatively similar polarization resistance. Yet, the contrast within tissue destruction is clear for cobalt-based materials. Depending on the implant application, careful consideration must be taken into account for any element that reacts adversely to the human body, since corrosion is also an atomic process and there is a risk of toxicity flowing into the bloodstream (Steinemann, 1998).

Stress shielding is a common phenomenon that takes place when stiffness of a metallic material exceeds the bone mechanical ability. This induces bone resorption which prevents bone growth in accordance with Wolff's law. One way to maneuver away from this issue is by designing a solid implant into the porous architecture (Carpenter et al., 2018). Titanium alloy can be fabricated into porous architecture especially for dental implants, while still maintaining its physiochemical properties thereafter, inducing deep bone integration in the implant (osseointegration) rather than just an interfacial bond. Consequently, avoiding stress shielding as a whole prevents future implant failure (Pałka and Pokrowiecki, 2018).

In addition, titanium is known to be dangerous for casting due to metal fuming and internal oxidation processes. Hence in the traditional way, it is fabricated and machined into the desired shape of an implant (Lütjering and Williams, 2007). Through 3D printing, the mechanical aspect of titanium (low modulus, high specific strength, and low density) can be preserved leading to biocompatibility (Matassi et al., 2013). However, pure titanium is normally avoided in bone implants for load bearing due to its low yield and tensile strength (Pałka and Pokrowiecki, 2018).

Implant failure can be due to several reasons: aseptic loosening, wear, and bacterial infection (Kong et al., 2018), as illustrated in Figure 3B. A study was conducted by Sailer et al. (2009) gathering the percentage of metal components undergoing failure or complication. It was reported that close to 50% was caused by fracture and 24% was due to corrosion. Therefore, emphasis on early surface treatment of implants is ideal to promote successful implantation to the biological system. For example, coating such as a titanium nanotube has an antimicrobial effect which promotes osteoblast formation on the implant surface (Wang and Tang, 2019). Since a titania surface exhibits bioinert behavior when surrounding a biological environment, the physical and chemical characteristics must be restructured. The formation of nanotubular structures promotes strong mechanical interlocking to bone cells as compared to a microstructure surface. As such, the key cellular activities progress while bacterial adherence declines. The inhibition of bacteria is crucial which can result in antibiotic resistance through biofilm formation. In addition, it is reported that through the fabrication process of specific nanotubes with diameters of 30 and 80 nm, a rough surface and low water contact angle are established, eventually aiding cell growth. Likewise, the chemical compositions of oxygen and fluorine are shown to induce both cell adherence and antibacterial ability (Peng et al., 2013).

Surface roughness, topography, wettability, and chemical composition are crucial to stimulate biocompatibility at the cellular level and interfacial bonding to the metal substrate (Tiainen et al., 2019). Surface roughness is an integral part of a material to allow bone and material interfacial bonding growth by interlocking cells to a material. It is widely known that a smooth surface provides poor stability to an implant as observed in Figure 5B. With surface roughness, osteoblast adsorption is observed which inhibits the osteoclast signal. As a result, full bone deposition is easily observed on the surface (Zhang et al., 2014). However, when it comes to high load-bearing application, surface roughness may pose an issue leading to high cycle fatigue. Osteolysis can occur when bonding strength within implanted bodies loosens thereafter leading to toxic ions leaching, especially when it comes to metallic elements such as cobalt (Wang et al., 2017).

Moreover, due to the lack of natural bone constituents such as bioglass and hydroxyapatite (HA) which form good chemical bonding to the bone, a metallic titanium implant in contrast has poor osseointegration and osteoinductive properties (Steinemann, 1998). It is also worth noting that despite bioceramic materials having similar bone constituents, it does not mean that they will form interfacial bonding with the bone. The key influence is heavily reliant on surface topography—the presence of surface pores hence allowing interlocking of bone cells (Davies, 2007). Similarly, a nonbonding biomaterial such as titania surfaces have the ability for bone bonding (Takatsuka et al., 1995). In hindsight, medical implants must be paired with both microstructure and macrostructure properties to achieve successful tissue integration. The surface chemical-wettability and surface energy of a contact angle below 90^° at the microlevel allow for better protein adsorption and subsequently cell attachment (Figure 5C). From the macro-level, the pore-related parameters provide pathways for vascularization and space for bone tissue growth (Song et al., 2019).

Despite 3D printing having multiple prospects, the primary use of 3D printing over traditional manufacturing methods, especially in the medical field, is the ability to print complex lattice structural cells in precise dimensions (Wang et al., 2017). It has been mentioned that a topological feature acts as a stress distributor influencing the mechanical capability by withstanding loads; in particular, the octet truss design has shown superior results compared to other cellular structures (Parthasarathy et al., 2011), whereas triple periodic minimal surfaces promote trabecular bone simulation (Fantini et al., 2017).

Porosity, pore interconnectivity, pore shape, and pore size are essential aspects for the functionality of a built implant, influencing the success rate of 3D printing in a biological system. Porous titanium alloy has a comparable Young's modulus to bone which does not trigger bone resorption as compared to other metallic implants. It is reported that poor bone ingrowth leads to implant loosening, henceforth tackling osseointegration bonding between bone tissue and an implant is vital (Wang et al., 2017; Pałka and Pokrowiecki, 2018). To achieve successful implantation, porous metallic implants should rejuvenate the function of the bone and promote regeneration of the damaged tissues. This is possible by establishing biocompatibility of materials with the living organism, adequate mechanical properties for load-bearing applications, and avoidance of stress shielding, since porosity influences compressive strength and elastic modulus which aid key cellular activities, such as cell adhesion, proliferation, and differentiation (Hrabe et al., 2013).

Although porosity is crucial, it is worth noting that bone ingrowth can be hindered without interconnected pore capability. With porosity, bone cells can penetrate, adhere, and encapsulate the pore structure to provide strong bone mechanisms (Takemoto et al., 2005). Adequate porosity with interconnecting pores allows formation of vascularization as a transport pathway for nutrient diffusion and metabolic waste, which help functional key cellular activities and tissue survival (Karande et al., 2004; Pałka and Pokrowiecki, 2018).

High surface-area-to-volume ratio, pore shape, and pore size are essential for cell attachment and growth. In a study by Mankani et al. (2001), it was observed that HA particles of 100–250 µm showed astounding bone growth, yet those smaller or larger than the critical size tended to show decreased bone formation. Interestingly, it was also noted that flat-sided particles showed no bone formation. It was also concluded that it was unclear how particle size and pore size could be interrelated, but it is hypothesized that the space between the particles will act as a gap indicating that pore size plays a role within bone formation. Subsequently, with the success of controlling printing design for pore sizes and shapes, 3D printing has provided flexibility for designers and engineers to adapt the mechanical behavior of bone tissue accordingly (Hrabe et al., 2013).

In an article by Rumpler et al. (2008), pores of high curvature (rounded corners) displayed more tissue growth as shown in Figure 6. The growth is reasoned by cell–cell neighboring interaction arising from mechanical forces, stimulating physical surface tension. High cell concentration is realized at the surface of high force. Although there is more work to be done to prove the research work, multiple articles have concluded or observed cell proliferation at the curvature as compared to other areas.

The commonly used surface treatment techniques for bio-implants such as micro-arc oxidation (MAO), laser surface texturing (LST), chemical etching, and alkali–heat treatment with HA electrochemical deposition are introduced in the Surface Treatment section, and a summary of their advantages and limitations is shown in Table 1, for each surface treatment method adopted.

Generally, bioceramic is not utilized for load-bearing applications due to its brittle mechanical behavior resulting from its ionic charge counterparts in contrast to metallic material. If implant failure occurs, bioceramic may pose catastrophic effects to the internal body. Despite the fact that the bulk of the implant is made from metallic material, the coating layer exhibits a bioceramic component which forms on the metallic surface. Bioceramic has great bone–tissue interaction which demonstrates biocompatibility in the biological system. Some common bioceramics are calcium phosphate, carbon nanomaterial, biphasic calcium phosphate, and bioglass. There are many advantages surrounding bioceramic coating as classified in Figure 9. Furthermore, bioceramic behaves as an excellent source as compared to other implant materials. It is devoid of both toxic ions such as cobalt or chromium and allergic reactions of nickel ions in stainless steel. It also minimizes disease transmission which is commonly observed in allogenic transplantation. Therefore, the intrinsic behavior of bioceramic has a pivotal role in advancing coating applications within the medical field.