1. Motivation
The utilization of three-dimensional (3D) printing technology for the fabrication of personalized 3D medical devices and tissue constructs represents a paradigm shift in healthcare, offering a myriad of possibilities that include replacing the once traditional one-size-fits-all approach in healthcare with a patient-centric model that is tailored to the unique physiology and requirements of each patient. The 3D printed medical devices are physical instruments or tools that can be used to diagnose, treat, or prevent medical conditions which include medical implants [
1], surgical models [
2], surgical instruments, and personal protective equipment [
3], whereas the 3D bioprinted tissue constructs can be fabricated using patients’ cells, are primarily used for repair and reconstruction of a recipient’s tissue [
4], [
5], [
6], [
7], [
8]. The 3D printing technology provides a versatile and highly automated platform for the fabrication of patient-specific 3D medical devices or tissue constructs with unprecedented precision and complexity. The key motivation behind the integration of 3D printing into the medical field includes the ability to design and fabricate bespoke 3D medical devices, implants, and prosthetics that seamlessly conform to a patient’s anatomy.
The global market for 3D printed medical devices is expected to grow from 2.7 billion USD in the year 2022 to 6.9 billion USD in the year 2028 at a compound annual growth rate of 17.1% [
9]. There is a high demand for 3D printed personalized medical devices or tissue constructs due to ① customization, ② increased complexity, and ③ print-on-demand ability. Notably, the development and production of 3D printed medical devices and tissue constructs involve complex considerations related to processes, materials, and regulatory requirements. These aspects are crucial for ensuring the safety, efficacy, and quality of the 3D printed parts and an in-depth understanding of these key considerations is vital for the development of safe, effective, and high-quality 3D medical devices and tissue constructs that meet the requirements of patients while complying with regulatory guidelines. This review article aims to highlight and discuss the various key considerations for the fabrication of 3D personalized medical devices and tissue constructs in terms of process, material, and regulatory considerations in the subsequent sections.
2. Publication landscape on 3D printing techniques
There are seven key 3D printing techniques based on the American Society for Testing and Materials (ASTM) standard classification (material extrusion, powder bed fusion, vat photopolymerization, material jetting, binder jetting, directed energy deposition (DED), and sheet lamination) as shown in
Fig. 1(a) and there is an increasing number of publications on 3D printed medical devices and tissue constructs over the last decade according to Web of Science using the following keywords: “3D printed” + products—“medical implants,” “surgical models,” “surgical instruments,” “personal protective equipment” or “tissue constructs” (
Fig. 1(b)). Further analysis on the 3D printing techniques for fabrication of 3D printed medical devices or tissue constructs was performed on Web of Science using the following keywords: “3D printed” + 3D printing technique (① material extrusion—“extrusion,” ② powder bed fusion—“selective laser melting,” “direct melt laser sintering,” “selective laser sintering” and “electron beam melting,” ③ vat photopolymerization-“stereolithography” and “digital light processing,” ④ material jetting—‘‘jetting,” ⑤ “binder jetting,” ⑥ “directed energy deposition” and ⑦ “sheet lamination”) + medical devices—“medical implants,” “surgical models,” “surgical instruments,” “personal protective equipment” or “tissue constructs.” Our preliminary results reflect the general publication landscape for 3D printed medical devices and tissue constructs; the most used 3D printing technique for fabrication of medical devices and tissue constructs is material extrusion (53.98%), followed by powder bed fusion (16.81%), vat photopolymerization (16.52%), material jetting (8.85%), binder jetting (2.36%), directed energy deposition (0.88%), and lastly sheet lamination (0.59%) (
Fig. 1(c)). All seven types of 3D printing techniques can be used for fabrication of 3D printed medical devices—material extrusion (41.83%), powder bed fusion (22.71%), vat photopolymerization (19.92%), material jetting (10.36%), binder jetting (3.19%), directed energy deposition (1.2%) and sheet lamination (0.8%) (
Fig. 1(d)), whereas only material extrusion (88.64%), vat photopolymerization (6.82%), and material jetting (4.55%) are suitable for fabrication of 3D bioprinted cell-laden tissue constructs (
Fig. 1(e)). It is important to note that further scientometric analyses are required for more detailed and in-depth investigations, which have been published elsewhere [
10]. In the following sections, a more comprehensive discussion of process considerations, material considerations, and regulatory considerations of various 3D printing techniques will be presented.
3. Process considerations
3.1. Different 3D printing processes
3.1.1. Material extrusion
Material extrusion is the most used 3D printing technique, and it is commonly used for fabrication of 3D printed medical devices such as prosthetics [
11], [
12], [
13], orthopedic implants [
14], [
15], [
16], surgical guides [
17], [
18], [
19], surgical models [
20], [
21], [
22], customized personal protective equipment [
23], [
24] and tissue-like constructs (e.g., bone [
25], [
26], [
27], [
28], cardiac [
29], [
30], [
31], [
32], and cartilage [
33], [
34], [
35]). Some of the commonly used materials include polymers-hydrogels [
36], [
37], [
38], [
39], [
40] or thermoplastics [
41], metals [
42], [
43], and even ceramics [
44], [
45]. In the material extrusion-based process, the material is selectively extruded through a nozzle or orifice to fabricate complex 3D printed parts in a layer-by-layer deposition approach [
46]. The material extrusion-based process can be categorized into two broad groups which involve ① material melting and deposition at high temperatures—fused deposition modeling (FDM)/fused filament fabrication/screw-assisted additive manufacturing (AM) [
47], [
48], ② pneumatic or mechanical-based material extrusion—direct ink writing and melt electrospinning writing [
49], [
50], [
51], [
52].
The material melt extrusion-based approach is dependent on numerous processing parameters which include temperature (nozzle, print bed, and chamber), nozzle speed, and slicing thickness [
53], [
54], [
55], [
56]. An optimal nozzle temperature (just below the material decomposition temperature) helps to enhance the weld strength of 3D printed parts; a study investigated the influence of nozzle temperature (210-250 °C) on the weld strength of acrylonitrile butadiene styrene (ABS) filament (has a range of decomposition temperatures beginning from 360 to 450 °C [
57]) and it reported that a maximum weld strength was obtained at the highest nozzle temperature (250 °C) [
58]. An optimal print bed temperature (dependent on the type of filament) facilitates improved adhesion of the 1st initial printed layer and minimizes the accumulation of thermal stress within the 3D printed parts [
59], [
60]. The optimal print bed temperature varies with the type of filament being used; a bed temperature of 115 °C was used for ABS filament [
59], whereas a bed temperature close to room temperature was used for both neat polypropylene and glass-reinforced polypropylene [
60]. The bed temperature is usually lower than the nozzle temperature to initiate the setting process for material solidification. Furthermore, an increase in the layer thickness can help to improve the elastic modulus [
61]. The dimensional accuracy of the printed parts in a material extrusion-based process is determined by various material- and process-related parameters such as material shrinkage [
62] and chamber temperature [
63]. The surface finishing of the final 3D material-extrusion printed parts can be improved by optimizing the process parameters [
64], the slicing strategy [
65], [
66], the part build orientation [
67], [
68], and chemical treatment [
69], [
70]. The surface finish, build time, and support structure are considered guidelines for deciding an optimal part orientation; it helps to minimize or eliminate excessive supporting structures and improves the overall surface finishing [
71].
Extrusion-based bioprinting is a 3D printing technology specifically designed for the fabrication of complex 3D tissue constructs using biological materials such as living cells and biomaterials. Some important process considerations during hydrogel extrusion-based bioprinting of tissue constructs include bio-ink printability [
72], shear stress [
73], [
74], and potential nozzle clogging [
75]. The bio-ink printability in extrusion-based printing can be represented by the loss tangent value (tan
—which can be defined as the ratio of loss modulus (
) to storage modulus (
) and printing within a suitable “printability window” is important for obtaining 3D tissue constructs with good shape fidelity. Next, the cell viability during the extrusion-based printing process is dependent on the shear stress experienced by the encapsulated cells within the extruded bio-inks, which is affected by the bio-ink viscosity, nozzle geometry, nozzle diameter, and printing pressure. There is also a potential issue of nozzle clogging when printing cell-laden bio-inks of high cell concentration using small nozzle diameter (∼100 µm). The optimal viscosity of a printable bio-ink is within 10-10
7 mPa∙s [
76] and a maximum concentration of ∼10
7 cells⋅mL
−1 (high cell-laden bio-inks may interfere with hydrogel formation) [
77], [
78]. The different types of cell-laden bio-inks used in extrusion-based bioprinting include alginate-based (5%-8% w/v) [
79], [
80], gelatin-based (5%-10% w/v) [
81], [
82], [
83], [
84], [
85], [
86], hyaluronic acid-based (1.5%-2.5% w/v) [
87], [
88], [
89], [
90], polyethylene glycol (PEG)-based (20% w/v) [
91], and Pluronic F-127 (≥ 20% w/v) [
92], [
93]. Furthermore, decellularized extracellular matrix (dECM) bio-inks derived from various tissues/organs are typically printed using extrusion-based bioprinting at 1%-4% w/v concentration [
94]. The dECM bio-inks are usually printed at a low temperature of ≤ 15 °C to mitigate pre-gelation and the temperature-dependent crosslinking occurs at 37 °C.
3.1.2. Powder bed fusion
Powder bed fusion is the 2nd most used 3D printing technique (16.81%); it can be used for the fabrication of 3D printed medical devices such as orthopedic implants [
95], [
96], [
97], dental implants [
98], [
99], [
100], and surgical instruments [
101], [
102], and a wide variety of powders (polymers, metals, and ceramics) can be processed [
103], [
104], [
105], [
106], [
107]. During the printing process, thermal energy is used to selectively fuse regions of a powder bed that are spread onto the building platform to obtain the 3D finished parts. The printing commences with orientating a pre-defined 3D computer-aided design (CAD) model within the build volume along with its supporting structures under inert conditions, scanning the scan path with a pre-defined combination of printing parameters, and lastly recoating with a fresh layer of powder particles [
103]. Examples of powder bed fusion processes include: ① selective laser melting (SLM), ② direct melt laser sintering (DMLS), ③ selective laser sintering (SLS), and ④ electron beam melting (EBM) [
103], [
104]. The key difference between SLM and DMLS is the temperature used to process the powders—the powders are fully melted at high temperatures in SLM to achieve a molten state, whereas the powder surfaces weld together during DMLS at a lower temperature [
105]. The laser-based powder bed fusion processes are usually performed in an inert environment, whereas the electron-based powder bed fusion process is performed within a vacuum chamber [
103]. Furthermore, each layer of powders in the EBM process is lightly sintered to minimize the build-up of electrostatic charges and repulsion of powder particles before subsequent scanning over the same toolpath to fuse the sintered powders. The excessive powders are removed via vacuum and further post-processing methods such as coating, sintering, or infiltration are performed.
Some critical process/material parameters for the powder bed fusion printing approach include energy input, powder recoating, coalescence, and cooling [
108]. The advantages of the powder bed fusion printing approach include the fabrication of complex structures with good surface finishing without the need for supporting structures (especially for polymer), whereas the limitations include low fabrication speed due to slow powder recoating speed and laser scanning speed. Some recent strategies to achieve 3D printed parts with improved density and surface finishing at a higher throughput rate involved mitigating the build-up of thermal stress and eliminating crack initiation and propagation via a re-melting approach [
109], [
110], [
111] and implementation of multi-laser scanning to improve the fabrication speed [
112], [
113], [
114].
3.1.3. Vat photopolymerization
Vat photopolymerization is the 3rd most used 3D printing technique (16.52%); it can be used for fabrication of 3D printed medical devices such as orthodontic devices [
115], [
116], [
117], hearing aids [
118] and surgical guides [
119], [
120], [
121], and tissue constructs (e.g., bone [
122], [
123], [
124], [
125], cartilage [
126], [
127], [
128], liver [
129], [
130], [
131], and neural tissues [
132], [
133], [
134], [
135]). The different materials used in vat photopolymerization include polymer resins [
136], [
137], metallic- [
138], [
139], [
140], and even ceramic- [
140], [
141], [
142] containing suspensions. The photo-curable resin is selectively crosslinked when a light source is projected onto the surface to initiate the free-radical photopolymerization process to obtain the completed 3D parts [
136]. The vat photopolymerization process can be classified into two main groups: stereolithography (SLA) and digital light processing (DLP); and it can be further categorized based on the position of the light source—bottom-up or top-down. The light source is positioned below the bottom of the vat through a window to crosslink the resin in the bottom-up SLA printing process, whereas the light source is positioned directly above the vat and building platform to crosslink the resin in the top-down SLA printing process. A digital micro-mirror device (DMD) is commonly used in the DLP process to facilitate rapid crosslinking of a single resin layer as compared to a single beam spot in SLA, which resulted in significantly faster fabrication speed for DLP [
143]. Notably, an optimal irradiation duration ensures sufficient macromer conversion to achieve strong interface bonding at the layer interface while minimizing over-exposure to prevent undesirable partial polymerization of surrounding resin. A recent study has demonstrated that the optimal UV exposure time for the fabrication of polydimethylsiloxane (PMDS) structures with a high aspect ratio was 2.5 s for each printed layer [
144]. A resin with a high molar extinction coefficient (
would facilitate crosslinking of a smaller voxel volume and mitigate over-curing. The light penetration depth can be increased by decreasing the photo-initiator (PI) concentration or reducing dyes/photo-absorbers. The addition of a non-reactive dye component or photo-absorber within the resin improves the printing resolution as it competes with the PI in light absorption. Some key considerations for the vat photopolymerization process include the laser source, printing parameters, vat design, resin viscosity, laser-resin interaction, and formulation of photo-curable resins (biocompatibility, solvent used, and the formation of free-radical by-products) [
136].
Unlike other bioprinting techniques (extrusion-based and jetting-based bioprinting) that deposit living cells via a nozzle, vat photopolymerization-based bioprinting uses biocompatible liquid bio-resins for the fabrication of intricate 3D cell-laden tissue constructs at high cellular density [
145]. Some important process considerations during vat photopolymerization-based printing of tissue constructs include the use of suitable laser wavelength [
146], [
147] and biocompatible PI [
148]. It has been reported that shorter wavelength is more detrimental to the living cells as the higher amount of energy leads to increased DNA cellular damage. Furthermore, an ideal PI should be highly hydrophilic (the presence of more hydrophilic groups decreases the cellular uptake of PIs and reduces the cytotoxicity effect on living cells) [
149], [
150] and crosslink at long wavelengths (formation of UV-initiated free radicals at short wavelengths has the most detrimental effect on the cell viability) [
148]. Significant progress has been made in the vat photopolymerization-based bioprinting technique and it is particularly valuable for the fabrication of complex 3D tissue constructs with high-resolution structures at high cell density. The different types of bio-inks used in vat photopolymerization-based bioprinting include alginate-based (5%-8% w/v), gelatin-based (2.5%-15% w/v) [
151], [
152], [
153], and PEG-based (10% w/v) [
154].
3.1.4. Material jetting
Material jetting is the 4th most used 3D printing technique (8.85%), and it can be used for the fabrication of 3D printed medical devices such as orthodontics devices [
155], [
156], [
157], anatomical models [
158], [
159], [
160], and tissue constructs (e.g., alveolar lung [
161], [
162], [
163], [
164], retina [
165], [
166], [
167], [
168], and skin [
169], [
170], [
171], [
172]). Different types of printable materials using the material jetting-based systems include polymers [
172], [
173], [
174], metallic- [
175], [
176], [
177], and even ceramic- [
178], [
179], [
180] containing suspensions. In the material jetting-based process, discrete droplets are selectively deposited onto an underlying substrate surface to fabricate the 3D printed part. There are different variants of material jetting-based systems which include inkjet-based (continuous or drop-on-demand (DOD)) [
181] and electrohydrodynamic jetting [
182]. During the printing process, the printhead moves horizontally to deposit the functional ink droplets (typically in the picoliter (pL) or nanoliter (nL) range). The applied pressure pulse from the actuator overcomes the surface tension of the ink to eject discrete ink droplets from the nozzle orifice. The deposited droplets on the substrate surface would coalesce and crosslink to form the desired shape in that layer, droplets for the subsequent layer are then deposited onto the preceding layer to eventually obtain the 3D finished parts.
The printability is influenced by the physical properties of ink (viscosity, surface tension, and density) and nozzle radius and it can be represented by the inverse of Ohnesorge number (Oh)—
Z value (defined as the ratio between the Reynolds number and square root of the Weber number) can be used to represent the ink printability in the material jetting-based process [
183]. The nozzle size has a significant influence on the printing resolution (∼1.2-2 times the nozzle diameter). Although the inkjet print-head is capable of printing at high printing frequencies (up to 30 kHz), it is recommended to deposit droplets at printing frequencies less than 500 Hz (due to inconsistent pressure within the print head at high printing frequencies) [
184]. As the nozzle diameter in most jetting-based printing systems is relatively small, it can only print ink with a narrow range of viscosities that range between 3 and 30 mPa·s [
185].
Jetting-based bioprinting is one of the pioneering 3D printing technologies to facilitate cell deposition for the fabrication of 3D tissues or organs. The various jetting-based bioprinting techniques include inkjet-based bioprinting [
186], laser-induced forward transfer (LIFT)/laser-assisted bioprinting [
187], acoustic-based bioprinting [
188], microvalve-based bioprinting [
189], [
190], [
191], and electrohydrodynamic jet bioprinting (also known as bio-electrospraying) [
192]. It has emerged as an important technique in the rapidly evolving landscape of biotechnology and regenerative medicine for manipulating, patterning, and assembling biologically relevant materials (cells, biomolecules, and biomaterials) in the form of ejected droplets to achieve specific biological functions in a DOD manner. Some important process considerations during jetting-based printing of tissue constructs include droplet impact velocity [
193], bio-ink viscosity [
194], shear stress [
195], and cell homogeneity [
196], [
197], [
198]. It was reported that the impact velocity of low-viscosity droplets (in the order of 1 mPa·s) has a significant effect on the cell viability and the cell damage increases exponentially with increasing droplet impact velocity as the shear stress increases strongly with droplet velocity. Furthermore, the study has also shown that a minimum droplet volume of 20 nL per spot helps to mitigate evaporation-induced cell damage [
193]. Conversely, an increase in bio-ink viscosity from higher polymer concentration resulted in improved cell viability even at higher droplet impact velocities due to the additional cushioning effect (higher energy dissipation) for the encapsulated cells during droplet impact on the substrate surface [
194]. Another work reported that even short-time exposure to high shear stress (> 5 kPa) is detrimental to the cell viability and its long-term proliferation profile during the jetting process [
195]. Another important consideration is the cell homogeneity within the bio-inks during the jetting-based printing process [
196]. Over time, the suspended cells within the bio-inks would sediment and adhere to the interior surface of the printing cartridge which lead to cell inhomogeneity. It was reported that the use of polyvinylpyrrolidone-based bio-ink helped to reduce cell adhesion and mitigate sedimentation during the printing process [
197], which improved cell homogeneity within the bio-inks. The different types of bio-inks used in jetting-based bioprinting include alginate-based (< 2% w/v) [
199], [
200], [
201], collagen-based (0.2%-0.3% w/v) [
202], [
203], [
204], [
205], gelatin-based (up to 4% w/v) [
206], [
207], and hyaluronic acid-based (up to 2.5% w/v) [
208], [
209].
3.1.5. Binder jetting
Binder jetting is the 5th most used 3D printing technique (2.36%), and it is used for the fabrication of 3D printed medical devices such as maxillofacial prostheses [
210], [
211], orthopedics implants [
212], [
213], [
214], and surgical models [
215], [
216]. The powder (polymers, metals, or ceramics—typical particle size of ≥ 30 µm) is the fundamental building unit in the binder jetting process and an ideal powder should have a uniform flow with negligible force between the particles [
217]. In the binder jetting process, a fresh layer of powdered material is spread into a layer and selectively joined into a pre-defined shape via selective deposition of organic liquid binder in a layer-by-layer manner to obtain the final “green” printed parts (generally brittle and highly-porous). Additional post-processing steps such as infiltration and/or sintering can be performed to enhance the mechanical properties after the removal of “green” printed parts. Some potential advantages of the binder jetting process include a wide range of powdered materials for the fabrication process, printing at room temperature and ambient air, no need for support structures, and a high fabrication rate, while the limitations include the need for post-processing steps (sintering and infiltration), potential distortion of printed parts during densification process, high surface roughness, and lower printing resolution [
218].
Some strategies to improve the powder flowability include the use of additives and/or powder coatings and dry powder [
219], [
220]. The choice of a suitable binder is important during the binder jetting process; it should have a low printable viscosity, consistent droplet dispensing, and clean burn-out characteristics [
221], [
222], [
223]. The binding process is highly sophisticated and dynamic; the nature of binder-powder interaction differs drastically with the wetting characteristics, geometry, diameter, and density of the powdered material and this interaction controls the geometry accuracy, mechanical properties, and final surface roughness of the printed parts. The droplet impact and binder infiltration are the leading causes of printing defects; high droplet impact may cause ejection of the fine powders from their original position (ballistic ejection) [
224] while slow binder infiltration leads to droplet coalescence on the powder-bed surface which interferes with powder spreading for subsequent layers [
225]. Some strategies to overcome these challenges include the addition of chemical fixatives to enhance the cohesiveness of the powder bed, improvement of the wetting characteristics of powder materials [
226], and reduction of the dispensed droplet diameter to increase the binder infiltration rate [
227].
A series of post-processing steps (debinding and sintering) are performed on the green body (metal) upon removal from the unbound powder. The debinding process removes the binder from the final printed parts before the sintering cycle and it is critical to minimize the residual stress during the debinding process. The different debinding approaches include thermal debinding, catalytic debinding, solvent debinding, and wick debinding [
228], [
229]. Next, sintering is performed to densify and strengthen the green parts via high-temperature heat treatment [
230]. The sinterability of the particles is dependent on their particle size [
231] and particle morphology [
232]. Material shrinkage is a common problem during the sintering process and this problem can be solved by using multimodal powder sizes [
233]. Additional post-processing methods can be performed to achieve near-to-full-density parts; these approaches are dependent on the type of porosity—closed or open. Hot isostatic pressing can be applied to significantly reduce or eliminate pores in printed parts with closed porosity [
234], whereas the infiltration approach is a process of filling a porous part with a liquid via capillary wetting. Some of the common infiltrants include epoxy or cyanoacrylate for polymer parts [
235] and bronze for metal parts [
236].
3.1.6. DED
DED is the 2nd least commonly used 3D printing technique (0.88%), and it is mainly used for the fabrication of 3D printed medical devices such as orthopedics implants [
237], [
238], [
239], [
240]. A broad range of metals and even ceramics can be utilized in the DED process; they include chromium, H13 tool steel, Inconel 625, stainless steels, titanium alloy and tungsten and so on [
241]. In the DED process, focused thermal energy (from lasers, electron beam (EB) [
242], or plasma/electric arcs [
243]) is used to melt the simultaneously deposited metallic material (powder or wire form) to create a pool of molten metal with an underlying heat-affected zone (HAZ) along the deposition path [
244]. The relative movement of the build plate to the deposition head is controlled by a computer numerical control (CNC) and the layer-by-layer fabrication process is repeated to obtain the finished part. A shielding gas is required to mitigate the oxidation reactions during the printing process by creating an inert atmosphere and a carrier gas is required to transport the metallic material through the deposition head to the pool of molten metal. The build plate (substrate) used in the typical DED process has a similar material composition as the material preform and the finished part is removed from the substrate after the printing process. Thermal monitoring using infrared cameras and/or pyrometers in the DED process is useful for data collection and feedback control [
244], [
245], [
246].
The key advantages of the DED process include the ability to fabricate functionally graded parts with different material/alloy concentrations and perform potential repair/cladding applications [
247], [
248]. The preform used in the DED process can be in the form of powder or wire. Both the powder-fed and wire-fed DED processes have their unique advantages and limitations. The use of the powder-fed DED process is more predominant as the blown powder dynamics can be controlled in real-time and can be altered to achieve the fabrication of complex structures with higher precision [
249], [
250]. Different methods such as gas atomization, water atomization, and plasma rotating electrode process (PREP) are typically used to produce the spherical powders (10-100 µm diameter) used in the DED process [
251]. The use of spherical-shaped powders helps to minimize the amount of inert gas trapped within the melt pool and reduce the overall porosity of the finished part. In contrast, the wire-fed DED process resulted in improved surface quality and process efficiency, but it requires a higher degree of control due to potential vibrations [
252].
Some key considerations during the DED process include using an optimal working distance [
253] and thermal distortion [
254], [
255]. The optimal working distance in the DED process is dependent on a few key parameters such as laser attenuation, distribution of powder concentration, heat, and mass transfer into the melt pool [
253]. The DED printing process produces high-temperature gradients which lead to residual stresses and plastic deformations that affect the structural performance and metallurgy of the printed parts. Thermal distortion of DED printed parts is influenced by the preheating and cooling conditions, printing parameters, and the geometry of the printed part [
256], [
257]. Particularly, depositing the first layer on a cold substrate during the DED process generates the most residual stresses due to high-temperature gradients, and reducing the substrate size can reduce the residual stresses due to lower heat flux and stiffness [
255].
3.1.7. Sheet lamination
Sheet lamination is the least commonly used 3D printing technique (0.59%), and it has only been used for the fabrication of 3D printed medical devices such as maxillofacial prostheses [
258]. It can utilize a wide range of materials such as polymers, metals, ceramics, and even composites [
259]. In the sheet lamination process, sheets of material of pre-defined shape/geometry are bonded together to form the 3D finished parts. The layer thickness is dependent on the thickness of the sheets of materials, and it determines the final quality of the printed part. There are different bonding mechanisms for the sheet lamination process which include gluing/adhesive bonding, thermal bonding, clamping, and ultrasonic welding. The gluing/adhesive bonding approach requires the use of adhesive-backed material sheets that provide bonding between adjacent layers. The thermal bonding approach facilitates the bonding of adjacent layers by heating the material slightly above its melting point in an inert environment. The clamping approach requires a bolt and/or clamping mechanism to hold the sheets together; this approach results in clamping forces that are perpendicular to the laminate interface and there is a possibility of delamination. The ultrasonic welding approach applies ultrasonic wave and mechanical pressure on the material sheets to facilitate diffusion-based bonding [
259]. The bonding/forming order can be categorized as “bond then form” or “form then bond” processes and the advantages and limitations of both processes will be discussed in the following sections.
The “bond then form” approach comprises sheet alignment, sheet bonding and lastly forming process (cutting according to the slice contour). A heated roller is typically used to melt the adhesives and bond the adjacent sheet layers (70-200 µm thickness) and a laser/mechanical cutter is used to obtain the defined shape/geometry of each layer. The remaining material sheet provides support for subsequent layers and the process is repeated to obtain the 3D finished parts. The advantages of the “bond then form” sheet lamination process include low material shrinkage, high fabrication speed, multi-material fabrication, and relatively lower material, machine, and process cost, whereas the limitations include the presence of glue leads to inhomogeneous mechanical and thermal properties, poor control over the dimensional accuracy due to inconsistent material sheet thickness, and challenging to achieve small internal structures [
260]. For the “form then bond” process, the defined shape/geometry is cut from the material sheet and then bonded with the preceding substrate until the 3D finished part is obtained. The “form then bond” process enables the fabrication of intricate features and channels, prevents cutting into previous layers, and eliminates the material removal step, but it requires an alignment system for accurate bonding between adjacent layers [
260].
We have provided a comprehensive discussion of process considerations, material considerations, and applications of various 3D printing techniques in the earlier sections.
Table 1 [
42⇓⇓-
45,
53⇓⇓-
56,
103⇓⇓⇓⇓-
108,
136⇓⇓⇓⇓⇓-
142,
172⇓⇓⇓⇓⇓⇓⇓-
180,
221⇓⇓⇓⇓⇓-
227,
253⇓⇓⇓-
257,
259⇓-
261] presents the analysis of different 3D printing techniques for fabrication of medical devices and tissue constructs.
ASA: acrylonitrile styrene acrylate; PC: polycarbonate; PEI/ULTEM: polyetherimide; PLA: polylactic acid; TPU: thermoplastic polyurethane; PEEK: polyether ether ketone; PA: polyamide; PCL: polycaprolactone; PLA: polylactic acid; PLLA: poly(L-lactic acid).
3.2. Build orientation and location
The build orientation and location of CAD models within the build volume of different AM processes play a critical role in influencing the properties of printed parts. The build orientation is denoted by the
X-,
Y-, and
Z- axes (order of letters is based on decreasing dimensions—longest to shortest) and the build location refers to the spatial position of the parted parts within the build volume (
Fig. 2(a)) [
262], [
264].
Selecting an optimal build orientation for a given part can improve the quality of the final printed part [
264], [
265], [
266]. Extensive studies were performed to determine the optimal build orientation and the optimization process leads to fewer support structures [
267], [
268], [
269], [
270], [
271], [
272], improved surface roughness [
267], [
269], [
272], [
273], [
274], reduced material cost [
267], [
272], [
273], reduced build time [
267], [
269], [
272], improved mechanical properties [
267], and lastly higher printing accuracy [
273]. The build orientation in FDM process (material extrusion AM process) influences the ductility and failure behavior of printed parts; the printed parts in on-edge (
XZY) and flat (
XYZ) orientation exhibited higher tensile strength, flexural strength, and stiffness as compared to parts printed in upright orientation (
ZXY) [
263]. Another study evaluated the influence of build orientation in the SLM process (powder bed fusion process); the printed Ti-6Al-4V parts in three different orientations (on-edge, flat, and up-right) showed similar α and prior-β grain sizes, similar elastic modulus regardless of its build orientation [
275]. However, the change in build orientation influences the ductility of printed samples, with the flat-orientated samples showing the lowest elongation at break values. This is likely due to the curling of the flat-orientated samples which led to the generation of defects during the processing of successive layers [
275]. Another study evaluated the tensile properties of PolyJet 3D printed parts (material jetting process) at different build orientations; the microstructure analysis showed that the direction of surface cracks/voids is dependent on the build orientation [
276]. The influence of build orientation on built parts from vat polymerization showed that the samples built in the upright orientation had the highest fracture load values [
277].
The mechanical properties of a printed part are directly affected by the build location; the variation in delivered energy at different locations (e.g., center vs edges) resulted in printed parts of different mechanical properties. A study that studied the influence of build position reported that the fabricated EBM parts showed finer microstructures at the top of the build part and coarser microstructures at the bottom of the build part [
278]. Furthermore, it was reported that the fracture toughness of fabricated AM parts at different locations varies significantly (
Fig. 2(b)) [
279]. Hence, both build orientation and location play a critical part in influencing the mechanical properties of printed parts.
3.3. Support structures
The presence of support structures offers more freedom of design to fabricate highly complex geometries using 3D printing technologies [
280]. The use of support structures is critical for the design and fabrication of complex structures; the number, type, and position of support structures used during the 3D printing process would affect the mechanical properties and geometric accuracy of the 3D printed parts [
281]. It is interesting to note that the critical overhang angle varies among the different 3D printing processes such as material extrusion, powder bed fusion, and vat polymerization [
281]. Automated algorithms are commonly used to determine the optimal number and position of supports to minimize material wastage and manual intervention may be required for more complex structures. Hence, it is important to analyze the geometry of printed parts (overhangs, high aspect ratio features, internal features, and thin features). Although the support structures can be removed via physical or chemical approaches, the removal process often leads to surface marks or residues.
A combination of optimized topology and build orientation for modified CAD models can help to minimize the required number of support structures [
282], [
283], [
284]. The use of topology optimization in the FDM process has significantly reduced the number of support structures; numerical experiments are validated through the FDM process to demonstrate the robustness and efficiency of the proposed support structure-constrained topology optimization method [
285]. Furthermore, a single-step optimization process can be implemented to simultaneously optimize the topology, build orientation, and support structures to minimize the number of support structures in the powder bed fusion process. High-quality solutions can be obtained using the proposed multi-optimization method (tested in a two-dimensional setting using a simplified support cost model) at a fraction of the time as compared to the fixed-orientation approaches [
286]. A two-step optimization process was performed using the MATLAB algorithm in another work to optimize the build orientation and cellular structures for minimal support structures [
287]. Another study reported the use of lattice support structures with a low volume fraction of up to 8% for Ti-6Al-4V parts built using DMLS to reduce the amount of material and time used in the fabrication of support structures [
288]. A similar study evaluated the influence of different lattice support structures on maximum normalized residual stresses in laser powder bed metal AM; the diagonal lattice support structure demonstrated the most significant stress minimization (
Fig. 3) [
289] (see
Table 2).
4. Material considerations
4.1. Material batch-to-batch consistency
It is critical to determine the consistency of the raw materials and final 3D printed parts as the printing process might induce some chemical and/or physical changes to the material. A series of steps to ensure material consistency include proper documentation of the material composition and evaluation tests for different AM techniques. Proper documentation of the printing material composition (raw material, additives, cross-linkers, or processing aids) includes recording the material supplier, material name and its Chemical Abstracts Service (CAS) number, material specifications, and material certificates of analysis (COAs). Furthermore, it is also important to perform appropriate characterization and aging tests based on the material type (metal, ceramic, and polymer) and form (powders, filaments, fluid, and hydrogels) used in each specific AM technique over time. The characterization tests for different material types include the evaluation of chemical composition for metal, metal alloy, or ceramic materials and the evaluation of chemical composition, molecular weight distribution, and temperatures (glass transition (T g), melting (T m), and crystallization (T c)) for polymers. The characterization tests for materials of the different forms include evaluation of powder diameter, powder size distribution and its rheological behavior, evaluation of filament diameter and its diametric tolerances, and evaluation of fluid viscosity or viscoelasticity and its pot life.
4.2. Material recycling
Material recycling in some of the 3D printing processes such as powders from binder jetting, DED, powder bed fusion processes, and resins from vat photopolymerization raised concerns over the possible alteration of material properties from its virgin state. The recycled free powder after a print may contain defects from elevated bed temperature and excessive sintering, whereas there might be partial polymerization of the resin during the photo-crosslinking process.
The recycled powders from the binder jetting process were stored and dried at 180 °C overnight in a vacuum to remove the moisture and passed through a 45 µm sieve. The presence of deformed powder particles was observed for unused stainless steel powders; the ratio of larger particles (> 30 µm) to smaller particles (< 10 µm) increases after repeated recycling process (
Fig. 4(a)) [
290]. Interestingly, the final printed parts using both fresh and recycled powders exhibited comparable hardness and yield strength even though the density of sintered parts from fresh powder was ∼1.5% higher than the recycled powder (
Fig. 4(b)) [
290].
Another study reported that the recycled titanium powders from the powder bed fusion EBM process showed an increased number of finer particles with a lower degree of fluidity, whereas the recycled nickel alloy powders for the powder bed fusion SLM process showed an increased number of coarse particles and higher fluidity [
291]. Although similar microstructures were observed from both new and recycled SLM powders, an increased porosity in fabricated parts using recycled powders was observed. Comparable ultimate tensile strength and yield strength were observed for both new and recycled powders, while there was a decrease in the ductility and impact toughness for the recycled powders [
291].
A recent study evaluated the influence of recycling on the zirconia-based slurry for vat photopolymerization; agglomeration of zirconia was observed (
Fig. 4(c)) and it increased the slurry viscosity (
Fig. 4(d)). Furthermore, there was a significant decrease in density of sintered parts from > 99% in pristine samples to ∼90% in recycled samples. Although the sintered parts from both samples showed similar hardness, there was a significant decrease in flexure strength (
Fig. 4(e)) and Young’s modulus (
Fig. 4(f)) for the recycled samples. Hence, the ceramic-based slurry may not be suitable for material recycling in the vat photopolymerization 3D printing technique [
292]. To date, there are only limited investigations on the suitability of recycled materials in different 3D printing processes and it is important to conduct a series of more extensive tests to carefully evaluate the suitability of material recycling for different 3D printing processes.
4.3. Protein adsorption on the material interface of printed parts
The protein adsorption on the material interface plays a critical role in regulating cell adhesion on the material interface. Protein adsorption at a material interface refers to the process by which proteins in a biological fluid bind onto the surface of a material. The key mechanism behind protein adsorption at the material interface is due to the displacement of organized water molecules (lower entropy, higher free energy) by proteins [
293]. The water molecules would influence the protein conformation and expose binding sites for interaction with the material interface; hence affecting the type and strength of interactions between the protein and material interface. Furthermore, the protein adsorption onto a material interface is also influenced by several other factors such as its protein structure (primary, secondary, tertiary, and quaternary) [
294], isoelectric point [
295], solubility [
296], and wettability [
297]. The free binding sites on the adsorbed protein layer can then facilitate cell adhesion and proliferation, which may be beneficial to tissue integration and wound healing [
298]. In contrast, excessive protein adsorption can lead to material “fouling” which can impair the device's function and trigger adverse reactions in the body [
299].
Different material surface properties such as surface chemistry, wettability, and topography can affect the protein adsorption and cell adhesion at the material interface (
Fig. 5(a)) [
300]. Substantial research works on surface chemistry aim to create uniformly coated material surfaces with different chemical functionalities that can prevent non-specific protein adsorption and provide bioinert surfaces. The surface chemistry plays a critical role in influencing the type, conformation, and strength of protein adsorption (
Fig. 5(b)) [
301], [
302], [
303], [
304]. The amount, type, and conformation of the adsorbed proteins (such as collagen, vitronectin, fibronectin, and laminin) regulate integrin binding, which can trigger different signaling pathways that control various cellular functions. The cells begin to spread and stretch on the surface after adhering to the surface; the protein-mediated adhesive interactions serve as mechano-sensors that facilitate cell-ECM interactions [
305], [
306]. The most common strategy to functionalize the material surface is to apply various monolayer/multilayer polymer coatings [
307]. Although this method is effective in reducing protein adsorption and thrombocyte activation, some of the existing challenges include the development of homogeneous and long-term stable monolayer coatings on chemically inert surfaces [
308]. Surface wettability can be referred to as hydrophobicity/hydrophilicity and it is an important factor that affects protein adsorption. It is well-documented that protein unfolding and surface coverage are significantly higher on hydrophobic surfaces than on hydrophilic surfaces [
309]. The proteins are shielded from the hydrophilic surface due to the strong water-surface interactions on hydrophilic surfaces and are consequently desorbed. Lastly, the surface topography of a device can be manipulated to control cell adhesion via protein adsorption. The different surface topographical cues including surface roughness, curvature and size of surface features can induce changes in the protein’s structure and orientation which affect the cell adhesion at the material interface [
310], [
311]. A high-throughput and quantitative approach was implemented to investigate the influence of surface roughness, protein concentration, and protein type on the protein-surface interactions [
312]. The surface roughness has a significant effect on the adsorbed proteins; the study showed that an increase in surface roughness (from 15 to 30 nm) resulted in a significant increase in saturation uptake (up to 600%). Furthermore, the nanostructured surfaces were shown to promote the formation of protein aggregates within the nanometric pores with an aspect ratio greater than 0.4 (depending on the characteristics of each protein) [
312]. Another critical parameter for surface topography is surface curvature; different studies have reported that an increase in the local curvature would lead to decreased protein adsorption [
313], [
314]. Hence, a good understanding of the material surface properties is important for the development of advanced biomaterials to manipulate protein adsorption and regulate cell adhesion on material interfaces.
4.4. Biocompatibility of printed parts
A biocompatible material does not induce any adverse reaction when in contact with surrounding living tissues [
315]. The different factors that influence the biocompatibility of a printed product include the chemical composition [
316] and surface properties [
317]. The final finished products may undergo a range of post-processing procedures which can lead to a significant change in the surface chemistry and topography which in turn affects the biocompatibility [
317]. It is recommended to test the medical devices in their final finished form and composition after the various post-processing steps (sintering, coating, cleaning, and sterilization).
The establishment of ISO 10993-5 aims to standardize various existing procedures and specifies the requirements and guidelines for the
in-vitro cytotoxicity testing of devices [
318]. The goal of
in-vitro cytotoxicity testing is to evaluate the potential of a device to cause cell damage and ensure its safety for use in humans. ISO 10993-5 protocol involves the use of established cell lines (American type culture collection (ATCC) CCL-1, CCL-10, CCL-75, CCL-81, CCL-163, and CCL-171 endorsed by ISO experts) for preliminary phase of investigation, which includes accurate and reproducible quantitative measurement of cell metabolic functions, via cell-counting procedures [
319], DNA level assessment [
320], and 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) assay [
321].
4.5. Degradation properties of 3D printed parts
Most medical devices are fabricated using biocompatible materials (metals, ceramics, or polymers) and these materials are likely to produce harmless degradation by-products as the implanted medical devices degrade over time within the human body. It is important to note that the degradation profiles of metal-based, ceramic-based, or polymer-based medical devices can vary significantly. Metals and ceramics are highly resistant to corrosion and degradation under physiologic conditions, making them durable for long-term implantation [
322], [
323], [
324]. In contrast, the polymers undergo different forms of degradation such as hydrolysis, oxidation, and mechanical wear [
325], [
326]. Although there are strategies to minimize/isolate the release of degradation by-products from implanted medical devices to surrounding tissues via the use of medical coatings [
327], [
328], the human body has the innate ability to metabolize and remove some of these degradation by-products [
329]. Hence, the choice of suitable material for a specific medical device should depend on important factors such as intended application, biomechanical requirements, and the device’s expected lifespan.
5. Regulatory considerations
3D printing technology has gained significant traction in the field of medicine and healthcare; different regulatory bodies around the world such as the Food and Drug Administration (FDA) from the United States, the European Medicines Agency (EMA) from the European Union, and National Medical Products Administration (NMPA) from China have been actively involved in the regulation and approval of 3D printed medical devices. Some of the important milestones for FDA-approved 3D printing technology include the first FDA-approved 3D printed cranial implant (OsteoFab Patient-Specific Cranial Device received FDA clearance in February 2013) developed by Oxford Performance Materials (USA) and the first 3D printed titanium-based spinal device (CASCADIA Lateral Interbody System received its FDA 510(k) clearance and CE mark in January 2016) developed by K2M Group Holdings (acquired by Stryker in year 2018). To date, numerous 3D printing companies have received FDA 510(k) clearance for their 3D printed medical devices based on the findings published by the FDA department (
Fig. 6) [
330]. Since then, several studies demonstrated the implementation of 3D printed medical devices into human bodies [
331], [
332], [
333], [
334].
Despite the significant progress in the regulation of 3D printed medical devices, there are currently no FDA-approved or cleared 3D bioprinted tissues/organs. To date, there is a lack of standardization in the field of 3D bioprinting (in terms of technology, materials, and process), which severely limits its adoption for clinical medicine. Furthermore, the 3D bioprinted cell-laden tissues/organs pose several challenges to regulatory authorities due to their increased complexity as compared to conventional 3D printed medical devices. Although good practices such as documentation of the cell source, evaluation of cell viability and function, and maintenance of a high standard of sterility can be adopted during the entire bioprinting process, some of the remaining bottlenecks include the logistics and unknown long-term safety in human hosts. There are limited research facilities that have the expertise and technology to fabricate 3D bioprinted tissues/organs; hence the patient-derived cells and extracellular matrices would need to be transported to a specialized biofabrication facility for tissue/organ fabrication and maturation before sending the matured tissue/organ back to the clinical facilities. The logistics can be highly challenging as transportation of living tissues/organs is a highly specialized and time-sensitive process that requires careful coordination and adherence to strict protocols to ensure that the 3D bioprinted tissues/organs remain viable and safe for transplant. Nevertheless, a recent study has offered some glimpse of hope in a first-of-its-kind clinical trial—the first transplantation of autologous bioprinted ear implant in humans (developed by 3D Bio Therapeutics) was demonstrated in June 2022, which highlighted the potential of 3D bioprinting technology for translational medicine [
335].
Although there is currently no FDA-approved 3D bioprinted tissue construct, numerous 3D printing companies have received FDA 510(k) clearance for their 3D printed medical devices. Some important regulatory considerations of 3D printed medical devices include sterility and mechanical properties of printed parts. It is important to adhere to the strict guidelines and requirements for the sterility of medical devices to mitigate the potential occurrence of infections. Furthermore, many medical devices (orthopedic implants, prostheses, and surgical instruments) require specific mechanical properties to perform their intended functions effectively and mechanical failures can lead to serious consequences during use. In the subsequent sections, more detailed discussions of these key areas will be provided.
5.1. Sterility of printed parts
A sterile implant reduces the risk of potential infections and ensures the safety and effectiveness of the implantation process [
336], [
337]. Sterility refers to the absence of pathogenic microorganisms and it can be achieved through the use of different physical/chemical sterilization methods such as heat [
338], radiation [
339], and chemicals [
340], [
341], [
342]. The sterility assurance level (SAL) is a measure of the probability that a medical device is sterile and the accepted SAL for medical devices is 10
−6 (one in a million chance).
The autoclaving process is a heat sterilization process that generates steam under pressure to eliminate microorganisms. The effectiveness of the autoclaving process in eliminating microbes is dependent on the temperature, the contact duration, and the free circulation of pressurized steam [
338]. However, the autoclaving sterilization method is not suitable for devices that are sensitive to heat or moisture. Gamma radiation is a type of ionizing radiation that can destroy microorganisms by destroying their DNA. Although it is an effective sterilization for devices that are sensitive to heat and chemicals, the high-energy gamma radiation can damage the polymer molecules over time and affect their chemical and physical properties. Chemical sterilization uses chemicals such as hydrogen peroxide (H
2O
2) [
340], ethylene oxide (C
2H
4O) [
341], or glutaraldehyde [
342] to kill the microorganisms. The chemical sterilization process involves exposing the devices to the chemicals at controlled temperature, humidity, and duration for effective sterilization. Although it can be used to sterilize devices that are sensitive to heat and radiation, some of the chemicals may be toxic and additional steps are required to remove the residual chemicals on the devices. Hence, the selection of suitable sterilization techniques is dependent on the material properties of the printed devices. Notably, the FDA recently approved the first H
2O
2 sterilization system from Steris Healthcare (Ireland) in August 2023 for the sterilization of 3D printed medical devices.
5.2. Mechanical properties of printed parts
A critical aspect of medical device testing considerations includes mechanical testing for patient-specific AM parts [
343]. Upon completion of different post-processing, cleaning, and sterilization steps, the final finished AM parts should be subjected to similar mechanical testing as those fabricated using conventional manufacturing methods (tensile, compression, bend, fatigue, impact tests, etc). The final finished parts or representative test coupons should be used for recommended testing and more emphasis should be placed on structures containing voids, supports, or porous regions. The device type and application determine the type of tests to be conducted (such as ultimate strength, yield strength, modulus, creep, fatigue, and abrasive tests). Furthermore, it is important to perform proper documentation of all the post-processing steps (removal of residuals, heat treatment, and final machining) as the material properties and performance of final finished parts are affected by the post-processing steps [
344], [
345]. As highlighted in an earlier section, the build orientation and location have significant effects on the properties of printed parts. Hence, it is important to develop a stringent manufacturing process with thorough monitoring to ensure the batch-to-batch consistency of final finished parts.
6. Concluding remarks
3D printing has attracted increasing attention for the fabrication of medical devices and tissue constructs in recent years. There are seven classifications of 3D printing processes, and a good understanding of the different 3D printing processes, along with their advantages and limitations, is important for the selection of a suitable fabrication technique based on the intended application. Next, biomaterials play a critical role in the mechanical and biological properties of the 3D printed parts. Some of the critical material considerations include the consistency of batch-to-batch raw biomaterials, the effect of material recycling on the material properties, the influence of material surface properties on protein adsorption and cell adhesion at the material interface, and biocompatibility and degradation properties of 3D printed parts. Lastly, the 3D printed medical devices and tissue constructs should be subjected to stringent regulations in terms of sterility and mechanical properties to ensure their safety and efficacy. Hence, it is timely to provide a concise review that highlights the various important considerations (process, material, and regulatory) for 3D printed medical devices and tissue constructs. We envisioned that a good understanding of the various key considerations in 3D printing (process, material, and regulatory) is critical for the fabrication of improved patient-specific 3D printed medical devices and tissue constructs.
Acknowledgments
Wei Long Ng acknowledges the support provided by the NTU Presidential Postdoctoral Fellowship.
Compliance with ethics guidelines
Wei Long Ng, Jia An, and Chee Kai Chua declare that they have no conflict of interest or financial conflicts to disclose.
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