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Corresponding author at: Centre for Nanotechnology & Advanced Biomaterials (CeNTAB), ABCDE Innovation Centre, Tissue Engineering & Additive Manufacturing (TEAM) Lab, School of Chemical & Biotechnology, SASTRA Deemed University, Tirumalaisamudram, Thanjavur 613 401, Tamil Nadu, India.
Tissue-engineered nerve guidance conduits (NGCs) are a viable clinical alternative to autografts and allografts and have been widely used to treat peripheral nerve injuries (PNIs). Although these NGCs are successful to some extent, they cannot aid in native regeneration by improving native-equivalent neural innervation or regrowth. Further, NGCs exhibit longer recovery period and high cost limiting their clinical applications. Additive manufacturing (AM) could be an alternative to the existing drawbacks of the conventional NGCs fabrication methods. The emergence of the AM technique has offered ease for developing personalized three-dimensional (3D) neural constructs with intricate features and higher accuracy on a larger scale, replicating the native feature of nerve tissue. This review introduces the structural organization of peripheral nerves, the classification of PNI, and limitations in clinical and conventional nerve scaffold fabrication strategies. The principles and advantages of AM-based techniques, including the combinatorial approaches utilized for manufacturing 3D nerve conduits, are briefly summarized. This review also outlines the crucial parameters, such as the choice of printable biomaterials, 3D microstructural design/model, conductivity, permeability, degradation, mechanical property, and sterilization required to fabricate large-scale additive-manufactured NGCs successfully. Finally, the challenges and future directions toward fabricating the 3D-printed/bioprinted NGCs for clinical translation are also discussed.
]. In vertebrates, this system is classified as the central nervous system (CNS) and the peripheral nervous system (PNS). The central nervous system includes the brain and spinal cord, which function as the command center regulating the rest of the body. The peripheral nervous system includes 12 pairs of cranial nerves, 31 pairs of spinal nerves and their cell bodies, and neuromuscular junctions [
]. The prime function of PNS is to connect the CNS with the rest of our body by conducting impulses to regulate body homeostasis in response to physiological and environmental stimuli. PNS comprises Schwann cells, neurons, macrophages, fibroblasts, and longitudinally-oriented blood vessels [
]. The motor and sensory neurons are polarized cells whose cell bodies reside in the spinal cord and branch to form dendrites. They extend to the peripheral organs and tissues, forming long cytoplasmic axons that aid signal transduction. The signal conduction initiates in the axon hillock away from the cell body and travels toward the synapses connecting the target organs. The axons, partially enclosed by the myelin sheath, an insulating material produced by Schwann cells (SCs), enhance the signal transmission process. Each axon is surrounded by a connective tissue called endoneurium, which is bundled together to form a nerve fascicle. The perineurium, which surrounds the fascicles are innervated by blood vessels externally. The nerve fascicles are again bundled together to form a complete peripheral nerve surrounded by an outer connective tissue layer called epineurium (Fig. 1). [
Abnormalities in the structural orientation of nerves and biochemical and electrical imbalances in the neurons affect the locomotory activities and functional organs like muscles, liver, kidneys, heart, lungs, etc. The most common causes of peripheral nerve disorders include injuries due to an accident, sport, or fall, genetic and autoimmune disorders (e.g., Charcot-Marie-Tooth disorder, Huntington's disease, rheumatoid arthritis, Sjogren's syndrome), and other clinical conditions (e.g., diabetes, Guillain-Barre syndrome, carpal tunnel syndrome) [
]. An injured or degenerated nerve can only regenerate under specific conditions – (i) the distance between the proximal and distal end of an injured nerve should not exceed the critical nerve gap (> 1.5 cm for rats, > 3 cm for rabbits, > 4 cm for pigs/humans [
]) (A critical nerve gap is defined as a minimum distance of an injured nerve from its proximal and distal end, beyond which nerve regeneration does not occur without the support of nerve grafts); (ii) the presence of neurilemma and endoneurium surrounding the axons; (iii) intact nucleus in the cell body and (iv) the two cut ends should remain in the same plane of injury [
When an injury occurs in a healthy nerve, the plausible changes that could occur in the proximal and distal segments (Fig. 2) are as follows – (i) Proximal segment changes: Chromatolysis involving cell nucleus migration from the center to the periphery of the cell body and dispersion of Nissl bodies; Retrograde degeneration of myelin sheath and Schwann cells up to the first node of Ranvier. (ii) Distal segment changes: Fragmentation of the axonal membrane, Schwann cells, and myelin sheath up to the axon terminals; Release of chemotactic factors such as 5-hydroxytryptamine and histamine by the endoneurial layer recruits macrophages to phagocytize the fragmented membrane and myelin sheath leaving a few Schwann cells alone intact. Both the proximal and distal segment changes are together referred to as Wallerian degeneration. After degeneration, a few pseudopodia-like extensions grow from the proximal cut-end of the injured nerve, termed fibrils or axon sprouts. The Schwann cells align in tracts along the endoneurial layer called bands of Büngner, which enable the axon sprouts to protrude towards the distal cut-end. Later, the Schwann cells proliferate, differentiate, and re-myelinate, surrounding the regenerated axons with intact axon cell bodies [
]. The axon regeneration of a normal nerve occurs at a rate of 2-3 mm/day, whereas the regenerated axons grow slowly at a rate of 0.25 mm/day. However, this depends on the availability and transport of neurotropic factors and cytoskeletal materials (such as actin and tubulin) [
]. In 1941, Seddon had broadly classified peripheral nerve injuries (PNI) into three categories based on ultrastructural changes in the injured nerve – Neuropraxia, Axonotmesis, and Neurotmesis. In 1951, Sunderland had elaborately graded Seddon's classification of PNI, which helped the surgeons to decide on nerve repair (Table 1) [
Peripheral nerve treatments involve either non-surgical or surgical methods depending upon the severity of the injury. Non-surgical methods, including drugs, physical exercises, and psychological treatments, may favor short-gap and minor injuries [
]. Complete nerve regeneration with functional and motor recovery may not be possible for complex and large nerve gaps, necessitating surgical interventions via direct suturing, nerve grafts, and nerve guidance conduits [
]. Direct suturing or neurorrhaphy allows the coaptation of nerves at proximal and distal ends using sutures or glue/sealants, which are suitable for small-sized injuries. In the case of medium and large nerve gaps, excessive tension is created during coaptation, leading to improper axon regeneration and thus requiring nerve grafting procedures. A recent survey has reported an increase in the demand for peripheral nerve grafts, with an expected compound annual growth rate (CAGR) of 7.65% from 2022 to 2030[
], which signifies the market potential and clinical demands for nerve grafts. Autograft is a gold-standard treatment used frequently for bridging large nerve gaps due to high clinical success rates and outcomes. However, the favorable outcomes of autologous nerve grafts are determined by the dimension and location of the injured nerve and severity of donor site morbidity [
]. On the other hand, nerve allografts have clinical limitations due to tumor formation, opportunistic infections and immune rejections post-implantation, requiring long-term immunosuppressive and anti-cancer drugs [
]. Tissue engineering plays an important role in the development of implantable neural support matrices like nerve guidance conduits (NGCs), nerve wraps, and connectors. These are developed via conventional scaffold fabrication methods such as solvent casting, particulate leaching, cast molding, electrospinning, and freeze-thaw technique using biocompatible materials (natural/synthetic biopolymers, growth factors, bioactive peptides, electro / magneto-active components) [
]. These neural support matrices are fabricated with simple architectures (such as hollow cylindrical tubes) connecting the proximal and distal ends of the injured nerves. They provide significant clinical advantages such as support guidance for regenerating axons, reduced infiltration of non-neural cells (e.g., myofibroblasts), and less scar formation [
Pierucci, A.; De Duek, E. A. R.; De Oliveira, A. L. R. Peripheral nerve regeneration through biodegradable conduits prepared using solvent evaporation. https://home.liebertpub.com/tea 2008, 14, 595–606.
Biswas, D. P.; Tran, P. A.; Tallon, C.; et al. Combining mechanical foaming and thermally induced phase separation to generate chitosan scaffolds for soft tissue engineering. https://doi.org/10.1080/09205063.2016.1262164 2016, 28, 207–226.
]. Other complications include increased sutures with a slow healing rate, causing patient discomfort post-surgery. Incorporating hydrogels, aligned fibers, fillers (e.g., growth factors, peptides, cell adhesion motifs), intra-luminal channels, and surface micro/nano-patterning in NGCs may facilitate deeper infiltration of cells and nutrients, which will eventually increase the neural cell alignment with required vascularization [
]. Though several potential designs are incorporated in NGC manufacture, only a few NGCs have received FDA clearance and are commercialized as Class II implantable devices (Table 3). Most NGCs were made of bovine/porcine-based collagen conduits and small intestinal submucosa-based xenografts, while a few NGCs were manufactured using synthetic-based biomaterials (e.g., polylactic acid (PLA), polyglycolic acid (PGA), poly (lactic-co-glycolic acid) (PLGA), polyvinyl alcohol (PVA)). However, regeneration with complete motor and functional recovery similar to the native nerves has not yet been achieved with the commercially available NGCs. These conduits have shown positive clinical signs post-surgery, predominantly for digital and non-critical nerve defects up to 3 cm. [
In recent years, additive manufacturing (AM) has emerged as an alternative to conventional scaffold fabrication techniques as it is capable of creating three-dimensional (3D) structures with complex internal and external microarchitectures of target tissues in a layer-by-layer approach [
]. This process requires a computer-aided 3D design/model generated from 3D modeling software/micro-computed tomography (µ-CT)/magnetic resonance imaging (MRI), fed into a 3D printer/bioprinter system which will be fabricated into 3D constructs as per the digital input [
]. Some of the remarkable advantages of AM technology in personalized medicines, development of tissue constructs, and drug delivery devices are (i) flexibility in selecting the printable bioinks, including cells, growth factors, polymers, and nanoparticles, (ii) ease of controlling the design and 3D printing parameters and (iii) high-speed production of the printed products with accurate dimensions and a high degree of reproducibility [
]. Hence, 3D-printed nerve conduits will have better features than conventional tissue-engineered scaffolds and other acellular allografts.
This review briefly discussed AM-based techniques, including their principles, strengths, limitations, and combinatory approach with conventional scaffold fabrication methods for developing 3D peripheral nerve tissue constructs. We have also summarized the choice of printable biomaterials and the physiological properties (e.g., structural design, mechanical strength, degradation, suture ability, and conductivity) required to develop an ideal 3D-printed NGC. Finally, the clinical challenges and regulatory concerns were outlined with possible solutions.
2. Additive manufacturing techniques to develop PNCs
Several conventional scaffold fabrication methods (such as electrospinning, solvent casting, freeze drying, gas foaming, and phase separation) are available to develop peripheral nerve conduits. Yet, developing patient-specific branched or unbranched conduits with native scale dimension and resolution remains challenging. As an alternative, additive manufacturing (AM)/rapid prototyping (RP) methods have emerged, offering better features in printed constructs, which are otherwise impossible with conventional methods [
]. According to the American Society for Testing and Materials (ASTM) group (ASTM F42 – Additive Manufacturing), AM or 3D printing can be classified into seven categories – vat photopolymerization, material jetting, binder jetting, material extrusion, powder bed fusion, sheet lamination, and directed energy deposition [
]. Among these categories, vat photopolymerization, material jetting, and material extrusion are the most widely used methods to fabricate cell-laden or cell-free peripheral nerve constructs. Each of these techniques requires a 3D digital model of a nerve conduit, which can be obtained from 3D modeling software (e.g., Solid works, CAD) or imaging techniques (e.g., µ-CT, MRI). These 3D models (eg. .stl, .stp, .max, .x3d, .vrml, .3mf, .obj, .fbx or .dae format) will then be converted into a printable language (.gcode file format) and are imported into computer-controlled rapid prototyping machines for the printing process [
] were also widely implemented techniques to fabricate next-generation NGCs. The available techniques used to fabricate the 3D printed / bioprinted NGCs are shown in Fig. 3 and are summarized in the following subsections.
2.1 VAT Polymerization
Vat photopolymerization is an additive manufacturing technique that involves the solidification of resin-based photopolymers upon selective light irradiation to produce 3D prototypes in a layer-by-layer fashion [
]. The photopolymers mainly consist of photoinitiators, monomers/oligomers, stabilizers, and diluents, which undergo polymerization (e.g., free-radical polymerization) by forming strong covalent linkages between the precursor photopolymers when exposed to light energy at a specific wavelength, causing changes in their physical and structural properties. For tissue engineering applications, acrylate and methacrylate functional group-tagged polymers were predominantly used as photoactive resins, owing to their ideal scaffold properties such as biocompatibility, bio absorbability, biodegradability and anti-immunogenic with desired physiological-mimic properties[
Stereolithography is the first patented and commercialized AM technique to create 3D prototypes for several applications, including biomedical sensors, heart valves, dental fillers, industrial assembly parts, and architectural prototypes for demonstration purposes. Most SLA-based 3D printers have an XY-axes movable low-power curing laser source (UV/visible wavelengths), a tank filled with photopolymer resin, a Z-axis movable receiver platform, and a computer interface to control the movement of the laser source and receiver (Fig. 3) [
]. The SLA process involves photopolymerizing the photosensitive resins by the lasers, which are focused at a single spot via deflection mirrors or digital micromirror arrays at selective regions into a predefined 3D pattern. When the first layer of the 3D model is developed, the receiver platform is lowered as per the layer thickness (typically ∼ 0.1 mm), followed by a resin coating and photopolymerization of the second layer. This process is repeated continuously until the 3D objects are completely developed. The main advantages of SLA are creating 3D objects layer-by-layer with high geometric accuracy and precision with smooth surfaces. This technique is suitable for developing cell-laden or cell-seeded constructs with higher cell viability and resolution using biocompatible photoactive polymers (resins). However, it is a relatively time-consuming and discontinuous process, requiring additional post-processing steps to cure the printed parts completely to obtain mechanically strong and stable 3D models [
]. Further, most SLA instruments widely utilize visible light lasers over UV lasers in biomedical applications as visible light wavelength does not cause cytoplasmic and genomic disruption in the printed cells. [
Recently, Li et al. developed three different SLA-printed nerve guidance conduits with grooved, hollow, and porous surface morphology in various dimensions (2–4 mm in diameter and 15–20 mm in length) at 250 μm printing resolution and 5 μm layer thickness. The conduits were successfully prepared by irradiating 405 nm violet light (50 mW) over the conductive photo-polymeric resin composed of polyurethane (PU), different concentrations of poly(ethylene glycol)-conjugated graphene oxide (0.5% / 1% / 3% / 5% of PEGylated-GO) (conductive component) and 5% 2,4,6-trimethyl benzoyl-diphenyl- phosphine oxide (TPO, photoinitiator). The cross-sectional SEM images of the 3D-printed conduits showed grooved and porous morphology, confirming the feasibility of the SLA technique to fabricate highly complex structures (Fig. 4A). Incorporation of the conductive component (5% PEGylated-GO) in the fabricated conduits showed increased hydrophilicity (∼72°), conductivity (1.1 × 10−3), and optimal mechanical strength (Young Modulus 2.52 ± 0.14 MPa; Tensile stress 3.51 ± 0.54 MPa) compared to non-PEGylated-GO conduits. However, these nerve conduits require detailed in vitro and in vivo assessments to prove the feasibility of SLA-printed conduits in repairing PNI [
]. Singh et al., fabricated an acellular hollow and multi-lumen polycaprolactone (PCL) conduit filled with aligned chitosan/gelatin cryogel-filled NGCs using the visible light-based projection SLA system. The polymeric resin was initially synthesized by premixing the PCL oligomers with the 1% camphorquinone (visible light photoinitiator), 1% (w/w) ethyl 4-dimethyl amino benzoate (reaction accelerator), and 0.2% (w/w) Orasol orange G (dye to control the penetration depth). Hollow nerve conduits were designed with a length of 1.9 cm, a wall thickness of 0.35 mm, and an inner lumen diameter of 1.5 cm, whereas multi-lumen nerve conduits were designed with a length of 1.9 cm, a wall thickness of 0.35 mm, and an inner lumen diameter of 1.5 cm with 4 porous channels of diameter 0.5 cm. Both the conduits had a sleeve length of 2 mm to aid the native nerves properly inserted within the conduit during implantation. Visible light lasers of wavelength 400 – 500 nm were irradiated at an intensity of 5.6 mW/cm2 over PCL-based resin to fabricate the nerve conduits. Scanning electron microscopic images confirmed the hollow and multi-lumen structures in the fabricated conduits, which also had a smooth surface and matched the dimensions of the 3D model designed using the software (Fig. 4B). In addition, the developed conduits were incorporated with biocompatible and neural cell-adhesive pre-gel precursors (2% low-viscous chitosan and 6.4% gelatin) and a crosslinker (1.5% glutaraldehyde) solution to support the axonal regeneration and Schwann cell proliferation. These pre-gel-filled conduits were subjected to a unidirectional freezing process using cold vapors to obtain aligned cryogel-filled PCL (aCG) conduits and a rapid freezing process at -20°C for 12-15 h to obtain random cryogel-filled PCL (rCG) conduits. Incorporating aCG cryogels inside the PCL conduits enhanced the tensile modulus (180 ± 10 kPa) rather than rCG conduits (140.03 ± 4.73 kPa), as aligned cryogels provided a unidirectional porous architecture to resist deformation upon the applied stress. In vitro culture of Neuro2a neuroblastoma cells in the aligned (aCG) / random (rCG) cryogel-incorporated nerve conduits for seven days exhibited a uniform distribution of cells with enhanced proliferation and expression of cytoskeletal proteins (actin) throughout the conduit compared to non-cryogel incorporated conduits. In addition, aCG cryogel-incorporated conduits displayed more cellular infiltration inside the conduits than rCG cryogel-incorporated conduits, indicating the presence of porous and aligned structures within the NGCs for efficient peripheral nerve regeneration. [
]. Evangelista et al., developed SLA-based single and multi-lumen nerve conduits using 30 wt% of high molecular weight poly (ethylene glycol) diacrylate (PEGDA) as polymeric resin, 0.5 wt% Irgacure 2959 as photoinitiator and RGDS-conjugated poly (ethylene glycol) (PEG) as a cell-binding component. The UV lasers were scanned at a speed of 205.36 mm/s with an intensity of ∼19 mW over the PEGDA precursor polymer for developing four single-lumen (outer length - 16.09 mm; inner length - 10.41 mm and outer lumen diameter - 3.97 mm; inner lumen diameter - 1.36 mm) and multi-lumen conduits (outer length - 16.09 mm; inner length - 10.41 mm; outer lumen diameter - 3.97 mm and inner lumen diameter - 0.51 mm) in a single run. After removing the unreacted PEGDA polymer and photoinitiator completely by continuous washing with deionized (DI) water, the conduits were allowed to swell overnight, lyophilized, and sterilized before in vivo implantation. The developed single and multi-lumen conduits were implanted successfully in Sprague Dawley (SD) rats with a 10 mm sciatic nerve gap. The authors have evaluated the regenerative efficacy of nerves after 5 weeks using histomorphometric analysis. Macroscopic observation of the implanted conduits after 5 weeks demonstrated faster degradation (i.e., disappearance) of single-lumen conduits compared to multi-lumen conduits. Partial regeneration of nerves was observed mostly in proximal and middle sections of the single lumen conduits with increased axon number and myelin sheath thickness equivalent to native nerves. However, very few Schwann cells were present in the distal ends of the single-lumen conduits. Further, multi-lumen conduits exhibited poor nerve regrowth on macroscopic and microscopic observations and were slowly biodegradable. These conduits also hinder nutrient availability and limit the gaseous exchange inside the conduits due to very small lumen diameter, surface topography, and poor choice of biomaterials. These disadvantages restrict their use for nerve regeneration applications [
]. The major components of a DLP system include movable XY-axes, a UV/visible light source, a resin tank, a digital light projector screen, a Z-axis movable receiver plate, and a computer-controlled program to regulate the axis movement (Fig. 3). The main advantage of DLP over the SLA process is the fast production of 3D structures with reduced printing duration, as the laser source in the SLA system crosslinks the resin at every single laser spot [
]. Similar to SLA, DLP also possesses a few disadvantages, such as limited choice of photopolymerizable biomaterial resins, strong odor due to the polymerization between acrylate groups and a photoinitiator, and more wastage of resins that result in high cost of the printed part [
]. Chen et al., developed a DLP-based biocompatible hollow peripheral nerve conduit using photocurable water-based polyurethane (PU) resin mixed with an oxidizing agent, polydopamine (PDA), and a cell guiding component, decellularized extracellular matrix (dECM). Initially, homogenous suspension of dehydrated water-based PU, 1.5% 2,4,6-trimethyl benzoyl diphenylphosphine oxide (TPO), 0.1% 2-hydroxy-4-methoxy benzophenone-5-sulfonic acid hydrate, 0.01% 4-hydroxy-2,2,6,6-tetramethyl piperidinooxy (TEMPO) and 30% 2-hydroxyl ethyl methacrylate (HEMA) were mixed with PU to prepare a photocurable PU vat resin. In addition, 2 mg/mL of dopamine, 1.2 mg/mL of ammonium persulfate, and freshly extracted dECM were added to the vat, as polydopamine enhances the blue-light absorbing capability and dECM supports neural cell proliferation and differentiation within the conduit. The conduit was designed using SolidWorks software with a length of 14 mm with four 0.7 mm circular holes (aid in suturing with native nerves) and an inner and outer diameter of 2 mm and 2.5 mm, respectively, fabricated by exposing blue light for 20 s at 100 µm per layer (Fig. 4C). The as-printed conduits were washed with ethanol and finally post-cured with blue light to obtain the PU/PDA/dECM nerve conduit. The printed conduits were evaluated for physicochemical, mechanical, degradation, and cytocompatibility assessments. The water contact angle of the PU/PDA/dECM conduits showed higher hydrophilicity (42.9 ± 1.8°) than PU and PU/PDA conduits, which may influence cellular behaviors. The addition of PDA and dECM did not affect the structural composition and printability of PU resin. The ultimate tensile strength of the PU/PDA/dECM nerve conduit (38.8 ± 1.6 MPa) was similar to the human nerves compared to the PU and PU/PDA conduits, which may be sufficient enough for surgical handling and implantation purposes. In addition, the biodegradation of the PU/PDA/dECM conduits immersed in simulated body fluid (SMF) exhibited faster degradation (7%) while maintaining structural integrity. Indirect in vitro cytotoxicity analysis of the PU/PDA/dECM conduits performed as per ISO 10993-12 standard using L929 fibroblasts displayed similar metabolic activity with the control (2D cultured) groups. In vitro culture of primary human Schwann cells (HSCs) for 7 days with the developed conduit showed more adhered cells with normal morphology, higher proliferation ability, and increased expression of HSC markers such as nestin, β3-tubulin, and microtubule-associated protein 2 (MAP2) compared to PU and PU/PDA groups. Thus, the developed DLP-based PU conduits were cytocompatible and supported the peripheral nerve regeneration [
Ye et al., designed and fabricated a multi-channeled gelatin methacrylate (GelMA)-based NGCs using a commercial DLP 3D printer. The conduit was successfully developed with a length of 5 mm, outer diameter of 6 mm, and 1.2, 1.6, and 2.0 mm-sized 4-channel internal diameters by exposing the UV light at a wavelength of 405 nm with an intensity of 12 mW/cm2 over the vat filled with 13.3% methacrylate tagged gelatin and 0.25% lithium phenyl-2,4,6-trimethyl benzoyl phosphinate solution (Fig. 4D). Generally, the quality of the DLP-based printed structures was determined using layer thickness, light exposure duration, and light intensity. In this study, the light exposure duration was varied to obtain NGCs with interconnected channels with desired dimensions and sufficient mechanical strength equivalent to native tissues. Fabrication of conduits at minimal light exposure duration (< 20 s) and over-exposure of light (>50 s) developed mechanically unstable (collapsed easily) and more hardened conduits (more brittle), respectively. However, conduits fabricated at optimal exposure duration (∼35 s) resulted in flexible conduits without any deformities on compression. In addition, the DLP-based fabrication of five GelMA conduits was obtained at 15.5 minutes with high reproducibility. Microscopic images of the printed GelMA conduits were similar to that of the designed 3D model, confirming the ability of the DLP technique as a fast and accurate technique to develop complex nerve architectures [
Material extrusion (ME) is a type of AM method to develop 3D objects in which the materials (polymeric filaments/cell-laden or cell-free hydrogels) are loaded, liquified, and continuously ejected out of the print head through the nozzle and selectively deposited layer-by-layer as per the pre-generated path from the CAD design. Most thermoplastic materials, shear-thinning polymeric extrudates, metals, ceramics, and paste-like materials (prepared using solid powders and binders) are commonly used in this technique and have wide applications in prototype manufacturing for industrial and medical sectors [
]. The advantages of the ME process include solvent-free manufacturing of 3D parts using multiple materials, low-cost, user-friendly equipment, and high production volume with varied dimensions. This technique also has a few limitations, such as the need for support structures when printing branched or angled structures, rough-surfaced final parts, and the nozzle radius limiting the size of the printed model. Depending on the material extruder, ME can be classified as plunger-based, filament-based or screw-based [
]. Techniques such as fused filament fabrication (FFF)/fused deposition modelling (FDM) and extrusion-based bioprinting fall under filament-based and plunger-based methods, respectively, which are used for developing cell-free or cell-laden nerve conduits. Several 3D printing parameters (geometry-based, process-based and structural-based), such as nozzle size, printing speed, filament melting temperature, layer thickness, etc., in these techniques need to be carefully optimized to achieve 3D printed NGCs with higher resolution. This is further discussed in detail in Section 3.2.
This method was first developed in the 1980s by S. Scott Crump, registered under the name "Fused Deposition Modelling". Stratasys Inc., commercialized several FDM-based 3D printers, referred to as plastic jet printing [
]. A typical FDM printer consists of a filament coil, a temperature-controlled extruder head, and a printing plate. Single/multiple polymeric filaments are melted and extruded via a nozzle from the extruder and printed over the plate layer-by-layer (Fig. 3). A few biocompatible thermoplastic polymers such as PLA, PU, and PCL were 3D printed to fabricate neural scaffolds in different architectures and evaluated the neuronal activities both in vitro and in vivo [
].sRodríguez-Sánchez et al., fabricated NGCs using 3D printed PCL membranes using a commercial FDM-based 3D printer. PCL filaments were melted at 80 °C and deposited at a path speed of 8.8 mm/s to obtain a two-layered square-shaped PCL membrane with the following dimensions: filament diameter – 396 ± 74 μm, thickness – 386 ± 41 μm, area – 225 mm[
], pore height - 312 ± 58 μm and pore length - 300 ± 51 μm. Later, the PCL membranes were rolled around 1.5 mm support and sealed using heating to obtain hollow PCL conduits with a 1.5 mm inner diameter and smooth outer surface (Fig. 5). The fabricated PCL-based NGCs were loaded with 1 × 106 canine multipotent mesenchymal stromal cells (AdMSCs) (isolated from adipose tissue) embedded in heterogenous fibrin biopolymer (composed of 50 μL cryoprecipitated water buffalo blood, 12.5 μL CaCl2, and 12.5 μL thrombin-like protein) hydrogel. The cell-loaded 3D-printed PCL NGCs and plain 3D-printed PCL NGCs were implanted successfully in a 12 mm sciatic nerve gap of female Wistar rats. After 12 weeks of implantation, the AdMSCs loaded 3D-printed PCL conduits showed better locomotive motor recovery (sciatic functional index (SFI): 65.12 and Tibial functional index (TFI): -72.69) compared to plain PCL conduits (SFI: -80.81 and TFI: -82.04). In addition, the thicker myelin sheath and more axon fibers were observed in the newly regenerated site of AdMSCs-loaded 3D-printed PCL conduits compared to the autograft group. Furthermore, the expression of neurotrophic factors (brain-derived neurotrophic factor (BDNF), glial cell line-derived neurotrophic factor (GDNF)), p75 neurotrophin receptor (p75NTR), Schwann cell marker (S-100), and neurofilaments were expressed equivalent to the autografts with increased intensity in the proximal regions [
Hsiao et al., reported the development of a 3D printed PLA construct with dimensions: length – 150 mm, width – 25 mm, and height – 0.3 mm. The construct was easily printed by melting the PLA filaments at a higher temperature (195 °C) with different gap widths (pore sizes 150 μm and 200 μm) within the scaffolds as per the 3D design. The top surface of the printed PLA 3D scaffolds showed an irregular surface with small rectangular crystals and hydrophobic behavior, which were confirmed using AFM images and water contact angle experiments, respectively. In vitro seeding of human dental pulp stem cells (HDPSCs) over the printed constructs displayed a steady increase in cell viability for up to 7 days. The 3D-printed PLA scaffolds were then coated with a thin layer of poly-L-lysine to induce the seeded HDPSCs to differentiate into the neural lineage. The poly-L-lysine-coated PLA scaffolds allowed the differentiation of HDPSCs when incubated with a neuronal induction medium containing BDNF, which was confirmed by the expression of neural markers such as glial fibrillary acidic protein (GFAP), MAP2, neurofilament-M (NF-M), nestin and β3-tubulin compared to non-coated PLA scaffolds [
]. Most commercial 3D printing filaments, such as TPU and PLA, are prepared at a high molecular weight or using aromatic raw materials. The degradation of 3D-printed constructs may take several years and release cytotoxic (e.g., acidic and aromatic compounds) byproducts. Hence, synthesizing polymers with tissue-specific-molecular weight may be beneficial to translate the developed 3D scaffolds clinically. Kaplan et al., synthesized a biocompatible polyester-based copolymer made of PLGA and poly-L-lactic acid (PLLA) with tunable biodegradable properties. They printed linearly oriented 3D construct using water-soluble butanediol vinyl alcohol (BVOH) as per the micro-CT data of intact spinal cord nearby the lesion site. PLLA/PLGA (1:1 ratio) solutions were injected into the highly-oriented printed BVOH construct and lyophilized to obtain a PLGA/PLLA construct with a uniform pore diameter of ∼ 240 μm. The aligned microporous channels in the PLGA/PLLA (7%) 3D scaffold possessed comparable mechanical strength to the native spinal cord and aided in the growth of more oriented axons throughout the printed scaffold [
Unlike the FDM process, this technique extrudes biomaterials (mostly polymeric solutions and bioactive component / cell-laden solutions) loaded in the print head, where filaments (or strands) are deposited layer-by-layer to develop the desired 3D structure (Fig. 3). It is classified based on the extrusion method as either pneumatic-based or mechanical force-based [
]. Pneumatic-based bioprinters use compressed air or N2, whereas mechanical-based bioprinters use mechanical force (direct force for piston-based extrusion bioprinters and rotational screw for screw-based extrusion bioprinters) to extrude biomaterial inks from the nozzle at a controlled flow rate and volume [
]. In addition, pneumatic-based bioprinters minimize microbial contamination due to the use of the sterile-filtered airway and allow smooth extrusion of liquid and gel-like bioinks. However, the devices used in the mechanical-based bioprinters cause cell damage as the extrusion devices provide more pressure and increase the risk of contamination though utilizing a low volume of inks [
]. Most extrusion-based NGCs were fabricated using pneumatic-based extrusion bioprinters using appropriate bioinks with shear-thinning and instant crosslinking properties (enzymatic, chemical, or physical methods) with better resolution and shape fidelity. Li et al., aimed to build a 3D neural construct composed of Schwann cells embedded in alginate-gelatin hydrogel using the extrusion bioprinting technique. The print head of the extrusion bioprinter (1 mL syringe) containing 2% alginate, 10% gelatin, and 2 × 106 RSC96 cells/mL was allowed to extrude to print into different shapes (square, round, and butterfly-shaped) at a thickness of 1 mm. The optimized printing temperature was 37°C with a chamber temperature of 8°C, printing speed of 0.15 mL/s, scanning speed of 3.5 mL/ss, and nozzle diameter of 25 G. The printed RSC96-laden constructs were crosslinked using 3% CaCl2 for about 3 minutes and the ability of the extruded 3D hydrogel towards cell viability, proliferation, and differentiation were evaluated. In vitro results after culturing for seven days showed uniform distribution of more live cells with increased proliferation compared to dead cells with no localized cell death within the printed construct. The RSC96 cells embedded in the 3D printed construct released more nerve growth factor (NGF) than the 2D culture, which was confirmed using ELISA assay. Further, the 3D printed constructs also displayed the neural proliferation marker expression (S100 marker) with no neural extensions (dendrites), possibly due to the absence of cell attachment motifs. This study demonstrated the feasibility of extrusion bioprinting in developing 3D neural constructs without affecting cell morphology and functionality [
Song et al., synthesized an electroconductive hydrogel (ECH) matrix composed of 5% gelatin methacrylate (GelMA), 1% PEGDA and in situ polymerized conductive matrix (0.2% chondroitin sulfate methacrylate (CSMA) / poly(3,4-ethylene dioxythiophene) (PEDOT) or 0.2% CSMA/PEDOT / tannic acid). The prepared precursor hydrogel solutions were loaded in the microextrusion-based 3D printer and extruded at a printing pressure of 50–90 kPa, needle size 25 G, printing speed 5 mm/s, bed temperature 10 °C, and printing temperature 22°C. The printed structures were crosslinked using blue light to obtain a stable multi-layered grid-shaped structure. Before printing, the pre-gel solution was evaluated for rheological and mechanical strength analysis. Results showed that the prepared ECH ink exhibited shear thinning ability and Young's modulus value (0.5–0.6 kPa) equivalent to that of the soft hydrogels. In addition, the pre-gel solutions exhibited conductive properties due to the faster charge transfer between the polyphenol structure in PEDOT chains through the dopant, tannic acid, which may enhance the bioelectrical signal between the adjacent host nerves and implanted scaffolds. The extruded 3D ECH scaffolds displayed better cell attachment and proliferation with extended neurite morphology when cultured with neural stem cells (NSCs) for 7 days, confirming the ability of extrusion bioprinting as a valuable tool for developing 3D nerve cell-laden constructs. [
Material jetting (MJ) is yet another promising technology in AM, which selectively deposits biocompatible materials onto the receiver platform to develop 3D objects with high dimensional accuracy with smooth surface [
]. Depending upon the material ejection process, MJ works in continuous or drop-on-demand (DOD) modes. Continuous mode-based MJ involves the ejection and deposition of materials continuously over the receiver. In contrast, the DOD-based MJ process involves the high-throughput ejection of materials in the form of droplets and deposits toward the platform as per the 3D design. The printing process was done layer-by-layer so that individual layers could crosslink or solidify by physical or chemical means, thereby creating a 3D construct with or without cells [
]. DOD-based methods such as piezoelectric/ acoustic wave / thermal-based ink jet, electrohydrodynamic jet, and laser-induced forward transfer (LIFT) are included in the material jetting process. Various biomaterials such as alginate, agarose, gelatin methacrylate, collagen, thrombin, fibrinogen, and polycaprolactone have been used to develop 3D tissue constructs using volume-based (1 to 7000 pL) and concentration gradient-based droplets [
Electrohydrodynamic (EHD) jet or e-jet printing, a maskless, non-contact and direct-write AM technique, involves the ejection of liquid (ink) by electrostatic (Maxwell's) forces from the nozzle tip towards the substrate with a printing resolution of 2-5 µm [
]. The e-jet system involves four main components: (i) A fluid supply unit that includes single or multiple syringes, a syringe holder and a flow controller for the continuous and constant flow of the ink from the syringe to the nozzle end; (ii) A high voltage supply unit to generate a strong electric potential to form stable cone jet. Exposure of the nozzle to the electric field develops the mobility of ions, causing their accumulation at the liquid surface. An increase in the electric field above its critical value induces deformation in the meniscus at the nozzle tip into a conical shape (Taylor cone). The solution is pulled off the nozzle tip towards the substrate (receiver). In addition, few EHD systems have pulse and function generators for applying pulsed charges to generate droplets from the jet (drop-on-demand process); (iii) Visualizing and imaging unit for viewing the Taylor cone formation and capturing the jet emission process and (iv) A receiver unit with the multi-axes movable, high-resolution platform for depositing the jet at a precise position (Fig. 3) [
]. Some of the shortcomings of conventional scaffold fabrication methods (e.g., electrospinning), such as uncontrollable porosity, pore size and interconnectivity of the scaffold, and lack of repeatability and customizability, may be improved by the EHD-jet method. Several modes of the EHD jet structures, such as micro-dripping, spindle, ramified-meniscus, stable/unstable cone-jet, oscillating-jet, multi-jet, and ramified-jet modes, can be identified at the nozzle tip while increasing the intensity of the electric field [
]. The continuous cone-jet mode is predominantly used to develop 3D-printed nerve conduits among different modes, as it produces high-resolution structures. Several research articles also emphasize that the EHD-jet technique has successfully fabricated 3D scaffolds for regenerative medicine applications. Nevertheless, this technique also possesses a few limitations during the fabrication process, such as difficulty in printing large and thick constructs, less availability of materials for the printing process, cell behaviors on e-jetted scaffolds, and the scaffold design [
Vijayavenkataraman et al., designed and fabricated a smooth, uniform pore-sized polycaprolactone (PCL) mesh-like scaffold using the 3D printing-assisted EHD-jet technique. The scaffolds were printed with varied pore sizes (125 ± 15 µm, 215 ± 15 µm, 300 ± 15 µm, 400 ± 15 µm, and 550 ± 15 µm) using the optimized solution and process parameters such as PCL concentration (70%), input voltage (2.4 kV), substrate speed (75 mm/min), constant flow rate (10 µL/min) and nozzle-to-substrate distance (2 mm). The printed PCL scaffold was then rolled and heat-sealed to form a tubular conduit with a diameter of 1.2 mm, length of 1–3 cm, wall thickness of ∼200 µm, and a greater porosity of greater than 60% (Fig. 6A). Mechanical strength and degradation analysis were conducted to determine the effect of pore size and porosity of the tubular 3D conduits. Results revealed that increasing the pore size and porosity causes a decrease in the mechanical properties (such as Young's modulus, yield stress, yield strain, ultimate stress, and ultimate strain) and increases the degradation rate, which may be because the scaffolds with large pore sizes and greater porosity allow better transfer of the surrounding medium, nutrients, gas, and growth factors. However, the mechanical properties are compromised as the pore sizes and porosities increase. In addition, ∼125 µm and ∼ 215 µm pore-sized 3D tubular scaffolds exhibited porosity and mechanical strength equivalent to the native peripheral nerves (porosity 60-80% and strength 6.5 to 11.7 MPa). In vitro culture of PC12 cells over the 3D-printed PCL scaffolds increased cell proliferation for up to 7 days in ∼125 µm pore-sized conduits compared to other pore-sized conduits. Further, gene expression of neural differentiation markers such as β3-tubulin, neurofilament–heavy chain, and GAP-43 was higher on ∼125 µm pore-sized PCL scaffold, which suggested the potential of the EHD-jet PCL tubular conduit to be efficient for treating peripheral nerve injuries[
]. Conductivity is an important property required for an ideal NGC, enabling better alignment/orientation, differentiation, and signal transmission for the growing axons. Hence, the same research group incorporated a conductive component, polypyrrole (PPy), at various concentrations (0.5%, 1%, and 2%) in 70% PCL polymer solution to develop a tubular PCL-PPy conduit using optimized EHD-jet printing conditions and evaluated its potential to treat nerve injuries. Incorporating PPy into the PCL scaffold did not alter the smoothness of the surface, chemical composition, and wettability properties. In addition, the conductivity and decomposition temperature of the PCL/PPy scaffolds were higher than the PCL scaffold. However, the printed features of PCL/PPY constructs were non-uniform and wavy due to the alteration in the viscosity of the PCL/PPy blended solution. Further, Young's modulus obtained from the stress-strain curve for PCL/PPy scaffold (35 ± 5.6 MPa for PCL/2% PPy) decreased and became softer compared to the PCL EHD-jetted group (204 ± 6.7 MPa), which may aid the neural cell growth and differentiation. Accelerated degradation of PCL/PPy scaffolds in alkaline conditions (pH ∼13) showed a faster degradation rate when compared to the PCL group. It also deteriorated the mechanical strength of the degraded scaffolds when increasing the PPy concentration, although the values are equivalent to human nerves (∼6.5 MPa). In vitro culture of human embryonic stem cells-derived neural crest stem cells (hESC-NCSCs) over the matrigel-coated printed PCL/PPy scaffolds had promoted growth and differentiation of stem cells into peripheral neurons, which was confirmed by MTS assay and RTPCR experiments, respectively. In addition, increased expression of neural markers such as β3 tubulin and neurofilament heavy chains were observed in the printed PCL/PPy scaffolds [
Inkjet bioprinting is generally a drop-on-demand (DOD), noncontact, low-temperature, and low-pressure micropatterning technique, which involves controlled deposition and positioning of uniform droplets (few pL with 20 to 100 μm resolution) from the ink cartridge towards the substrate placed at a minimum distance from the cartridge (usually below 1 mm) [
]. The inkjet 3D printer/bioprinter consists of a dispensing cartridge loaded with less viscous bioinks and a receiving substrate (culture plate or dish), which are software-controlled, moving in three axes to reproduce the digital pattern (Fig. 3). Depending on the droplet dispensing mechanism, this technique can be classified as continuous inkjet (CIJ) and DOD inkjet [
]. CIJ methods were not used in developing biological constructs as it creates sterility issues and uses electric or magnetic fields to generate a train of redundant droplets at higher frequencies. In contrast, the DOD inkjet method generates uniform-sized droplets at lower frequencies and achieves higher printing resolution. Based on droplet generation in DOD inkjet, they are further classified as thermal-based, piezoelectric actuator-based, and electrostatic-based [
]. Thermal-based inkjet printers/bioprinters use localized thermal energy (typically up to 300°C) to generate air bubbles inside the cartridge, causing the ejection of droplets from the bio ink-loaded cartridge through the nozzle. Piezoelectric-based inkjet printers/bioprinters employ voltage to the piezoelectric material connected to the nozzle, causing a sudden change in the cartridge volume and resulting in the formation of droplets [
]. Electrostatic inkjet printers eject droplets when the electric current is applied to a platen, causing expansion of the cartridge and deposition of the loaded materials. The deposition of high viscous inks using these printers would require increased deposition temperature / actuating current, which results in damage to the embedded biomolecules, clogging of the nozzle, and unreliable cell encapsulation, requiring an upper cut-off for the viscosity of the bioink [
Among these techniques, thermal-based and piezoelectric-based inkjet technologies have been widely used for bioprinting 3D constructs, as the localized temperature / piezoelectric actuating voltage in the cartridge do not cause any detrimental damage to the stability and functionality of the deposited biomolecules. For instance, Sun et al., have patterned the short self-assembling I3QGK peptide-based bioink over the regenerated silk fibroin (RSF) films using an inkjet printing technique to support the growth of neuronal cells. Initially, 1 mg/mL I3QGK peptides were dissolved in 20 mM HEPES buffer solution and optimized the nanofiber self-assembly process at different time points using atomic force microscopic (AFM) images. Results revealed that the self-assembly of I3QGK peptides started within 3 hours of incubation, demonstrating the aggregation of I3QGK molecules through hydrophobic interactions between the I3 tails and the formation of bilayers, followed by uniform nanoribbon formation with a length of 5-10 nm and width of 30 nm after incubation for two weeks. Using a piezoelectric DOD ink-jet printer, the I3QGK peptide nanofibers were deposited into multiple-layered line patterns (Fig. 6B) on the RSF/I3QGK peptide-coated (40 mg/mL, 8000 rpm, 25 s) glass substrate using the printing parameters: actuation voltage – 90 V, frequency – 300 Hz, nozzle size of print head/cartridge – 60 µm and distance between the print head-receiver – 1 mm. Strong electrostatic interactions between the negatively charged RSF and positively charged peptide solutions enabled more nano-fibrillar structures over the substrate with increased peptide concentrations. In addition, in vitro culture of PC12 neuronal cells at a density of 10,000 cells/cm2 for two days over the ink-jet deposited 5-layered RSK/I3QGK peptide substrate displayed cells with good adhesion property only on the peptide substrates with extended actin filaments. This study showed the feasibility of piezoelectric-based inkjet technology in printing biopeptides without affecting their functionality. [
]. The laser-based bioprinter system is primarily composed of three essential components connected to a software-controlled unit – (i) pulsed or continuous-wave lasers as the energy source; (ii) ribbon/donor usually made of a quartz/glass slide coated with or without laser energy absorbing/sacrificial layer followed by a micrometer-thick bioink layer and (iii) substrate for receiving the propelled bioinks in a predefined pattern (Fig. 3) [
]. Matrix-assisted pulsed laser evaporation–direct write (MAPLE-DW), laser-guided direct write (LGDW), absorbing film assisted–laser induced forward transfer (AFA-LIFT) and laser-induced backward transfer are some of the laser-based methods that are explored towards the bioprinting of tissue constructs. Generally, when the laser hits the donor, the coated sacrificial layer (or the part of the bioink layer) on the donor evaporates to form a vapor bubble, which expands and collapses due to pressure formed inside the bubble, causing falling or displacement of the bioink in the form of droplets towards the receiver [
]. Most laser bioprinting techniques utilize nano/femtosecond pulsed UV/IR/NIR lasers to deposit several cell types and biomaterial inks (e.g., polymers, proteins, ceramics) with higher cell viability (> 95%) and sustained growth and functionality [
]. Various factors, including the rheological properties of bioink, the coating thickness of bioink on the donor/substrate, laser pulse energy, and frequency, printing speed, and gap distance between the donor and substrate, influence the droplet resolution. Various reports have shown that by changing these parameters, the droplet resolution of less than 100 µm can be achieved. However, developing thicker 3D constructs (> 1-2 mm), scalability, and crosslinking feasibility are some downsides that need to be considered while printing 3D constructs using laser-based methods [
]. Tortorella et al., patterned arrays of laminin peptide over the receiver coated with biodegradable poly (lactic-co-glycolic acid) (PLGA) films. The laser bioprinting process was performed using a NIR-based Nd:YAG laser of wavelength 1064 nm with laser pulse duration of 10 ns, power of 0.2–1 W, and frequency of 5–50 kHz connected to a computer-controlled motorized unit and positioned the laminin droplets on the PLGA substrate at a gap-distance of 600 µm. Atomic force microscope images confirmed the aggregation and complete adsorption of printed laminin peptide over the PLGA film within 120 minutes, demonstrating the unaltered functional ability of the laminin peptides. Further, in vitro seeding and culture of neural stem cells (NE-4C) over the LAB-printed laminin / PLGA scaffolds allowed the cells to attach firmly, proliferate and differentiate as clusters with desired orientation along the laminin deposited regions, which were confirmed by SEM and fluorescent images (Fig. 6C). This study showed the feasibility of LAB technique to develop implantable neural scaffolds with aligned topographies [
Melt electrospinning writing (MEW) is a contactless, nozzle-based 3D printing technique that continuously ejects highly viscous or molten polymers, allowing ordered deposition of complex 3D geometrical structures with a micro-to nano-scale resolution [
]. Thus, it combines the principle of electrospinning and extrusion-based methods. Initially, the thermoplastic polymers (e.g., polycaprolactone, polyvinylidene fluoride, polypropylene) are melted and extruded as a spherical bubble through a nozzle by a pneumatic or volumetric-based dispenser. Applying a high voltage in the nozzle causes the droplet to form a Taylor cone due to electrostatic forces exerted between the nozzle and the collector. Thin stable polymer filaments are continuously drawn towards the collector from the Taylor cone, which depends upon several parameters such as voltage, dispensing pressure, polymer melting temperature, collector speed, the distance between the nozzle and collector, nozzle diameter, and polymer viscosity, thereby producing predefined 3D patterned scaffolds [
]. Unlike the conventional electrospinning process, this method is solvent-free and suitable for low-conductive polymers, preventing volatility, toxicity, and electrical instability issues. In addition, MEW can produce ultrathin polymer filaments (∼1 µm – 50 µm) with higher porosity (up to 90%) than other 3D printing techniques [
]. However, it limits its usage in developing large volume or thicker constructs due to the repulsive or semiconductive nature of the printed polymer filaments that result in distorted architectures, which can be prevented by coordinated adjustment of z-axis movement and the applied voltage [
]. Chen et al., have fabricated a grid-patterned PCL scaffold using the MEW technique, where the scaffolds were designed and printed with different inter-fiber spacings (100 µm, 200 µm, and 400 µm) and precise stacking of layers (2, 5, and 8), which was confirmed by SEM images (Fig. 7A). An increase in scaffold surface area is one of the requirements for faster regrowth of neural cells, which can be achieved by developing scaffolds with reduced spacing between the fibers. Hence, the diameter of the extruded PCL fibers was adjusted by optimizing the process parameters such as jet lag length, the collector and syringe speed, air pressure and the applied voltage. The jet lag length, the horizontal length between the contact point of the fiber on the collector and the central line directly below the nozzle, played an important role in achieving the precise stacking of PCL fibers to obtain 3D scaffolds. The decrease in diameter of the extruded PCL fibers was significant when increasing the lag length, collector speed and voltage as the fibers and the polymer melt got stretched out rapidly and easily. On the contrary, the diameter of the PCL fibers increased upon increasing the air pressure due to the increase in the extruded quantity of the polymer melt, producing deposition of thicker fibers. In addition, the mechanical property of the different layered and inter-fiber-spaced MEW-printed PCL scaffolds was identified to match the mechanical strength of the native nerves. An increase in the inter-strand spacing between the printed 3D scaffolds (100 µm, 200 µm, and 400 µm) significantly decreased the modulus values for five layered PCL scaffolds. However, there was no significant difference in the tensile modulus of different layered PCL scaffolds with 100 µm spacing. These results suggest the effectiveness of the MEW technique in creating highly-oriented tissue-engineered 3D scaffolds for treating peripheral nerve injuries [
The Kenzan or microneedle-based method, an updated version of extrusion-based bioprinting, is a biomaterial-independent bioprinting approach providing spatial and temporal control over the positioning of cell spheroids (loosely agglomerated cells) without the aid of biopolymeric solutions or hydrogels (as observed in traditional bioprinting methods) [
]. In this process, the homogenous/heterogenous cell spheroids are primarily formed and robotically positioned in the fine needle arrays (temporary support) as per the predefined 3D design (Fig. 3). The close contact between each spheroid supports the extracellular matrix formation, stabilizing the 3D structure during the in vitro culture conditions. Later, the temporary support is gently removed from the matured 3D tissue construct [
]. A few factors that need to be considered during the Kenzan-based bioprinting include (i) the size of the spheroid, which determines the inter-distance between the microneedles; (ii) spheroids formed via multiple cell types must be uniform and rearranged continuously. However, the spheroid size (e.g., 500 μm[
]) should be kept below the oxygen diffusion limit as increasing the size of the spheroid may develop hypoxic conditions due to a prolonged incubation period. The strong-adherent cells tend to occupy the core when compared to loose-adherent cells, which causes the core cells to starve for nutrients and oxygen and is predominantly observed in large-sized spheroids. The creation of a balanced environment for cell-cell interaction and extracellular matrix (ECM) formation within the spheroids are important for developing stable 3D constructs during the printing process [
]. To evaluate the feasibility of this technology towards clinical translation,. Mitsuzawa et al., have developed a scaffold-free hollow conduit using autologous dermal fibroblasts and examined the regenerative efficacy of the developed conduit by implanting it successfully in an ulnar nerve defect of large animals such as dogs. Initially, the fibroblast-containing spheroids were formed in the low adhesion 96-well plate in 24-48 h using canine dermal fibroblasts with a diameter of about 550 µm. The formed spheroids were aspirated using a fine suction nozzle and skewered to the circularly arranged micro-fine needles forming an 8 mm diameter tubular conduit (Fig. 7B). The formed tubular conduit was then cultured for one week to allow the fusion of spheroids in the temporary needle support, followed by the removal of the support. Later, the Bio-3D construct was placed inside a silicone conduit of 5 mm external diameter and cultured in a bioreactor for 20 days to attain the desired mechanical strength and function. The implantation of the developed Bio-3D conduits in a 5 mm ulnar nerve defect of dogs for 10 weeks showed complete regeneration with more axons with thick myelination both in the mid and distal regions of the conduit, which was comparable to the intact (normal) ulnar nerve group. In addition, immunohistochemical analysis in the mid-regions of the Bio 3D conduit group showed the expression of neural markers such as NF-200 and S-100, which confirmed the presence of neurofilaments and Schwann cells. However, in this study nerve autograft or FDA-approved nerve conduit group was not included for comparing the results with the developed Bio 3D conduit. Further, a 5 mm ulnar nerve gap did not match with the critical-sized nerve injury gap for large animal models and comprehensive motor and sensory nerve recovery evaluation needs to be conducted to evaluate the regenerative efficiency of the Bio 3D conduit comparable to commercial nerve conduits [
]. In another study, Yurie et al., developed a 8 mm long scaffold-free conduit using normal human dermal fibroblasts-based spheroids of ∼750 µm diameter and implanted in immune-deficient rats with 5 mm sciatic nerve defect. After eight weeks of implantation, the developed 3D implants (and silicone conduits) were excised and examined for peripheral nerve regeneration using histological and morphological measurements. Histological analysis revealed the presence of more myelinated axons with neural tissue formation compared to the silicone group. In addition, the immunohistochemical analysis resulted in the positive expression of NF-200 (neurofilaments marker) surrounded by the expression of S-100, confirming the presence of neural cells. Moreover, the newly formed axons were well-myelinated, surrounded by anterior tibialis muscle, with reduced atrophy in the 3D implant group compared to the silicone conduit group. Further, the functional, sensory, and motor recovery was evaluated using kinematic analysis (gait analysis), pinprick, and toe-spread tests. The pinprick and toe spread test showed similar results after eight weeks, with no significant difference between the 3D conduit and silicone groups. The gait of rats was evaluated in the treadmill, moved at a constant speed of 10 cm/s, which showed a higher angle of attack (AoA) with decreased plantar flexion angle of toes during the terminal swing phase in the 3D conduit group compared to the silicone group. These results indicated a greater sensory, motor, and functional recovery in the 3D conduit group [
Researchers have recently attempted to fabricate an ideal tissue-engineered scaffold (or biomedical product) with diverse properties by combining two or more scaffold fabrication techniques to provide an efficient treatment strategy and accelerate the regeneration process [
]. The combination of conventional scaffold fabrication with the 3D printing techniques allows the use of advantages from each technique, thereby masking the limitations of other techniques. This combinational strategy also enables the simultaneous use of different materials to develop multi-layered heterogenous 3D structures, which may impart adequate nerve-equivalent properties. For instance, Liu et al., have fabricated a peripheral nerve conduit by combining three different scaffold fabrication approaches such as electrohydrodynamic (EHD) jet printing, dip-coating, and electrospinning using two different biocompatible materials such as polycaprolactone (PCL) and gelatin. The triple-layered conduit was prepared in three steps - the inner layer was prepared using EHD jet printing of PCL polymer and rolled to form a hollow channel with an inner diameter of 1.5 mm and wall thickness of 1.2 mm; the middle layer was formed by dip-coating gelatin polymer, followed by crosslinking using microbial transglutaminase (mTG), and finally, the gelatin-coated conduit was wrapped with the electrospun PCL mat, forming the outer layer (Fig. 7C). The combination of different fabrication strategies in developing the conduit provides the following advantages - (i) the inner EHD-jet printed PCL polymer results in oriented fiber geometry with controlled pore size and porosity, enabling directed neural growth; (ii) the middle dip-coated gelatin layer over the EHD-jetted PCL matrix improves the cell adhesion and proliferation and (iii) the outer electrospun PCL fibers offer nanofibrous features with adequate mechanical strength and protect the developed conduit from fibroblast infiltration. The fabricated conduit had an adequate tensile strength (7.530 ± 0.151 MPa) compared to EHD jetted PCL conduit (2.709 ± 0.108 MPa), and EHD jetted PCL structure coated gelatin conduit (3.874 ± 0.135 MPa). In addition, the mechanical properties of the triple-layered conduits can be modulated and custom-made by varying the thickness of different layers such as EHD-jetted PCL matrix, gelatin-coated layer, and electrospun PCL mat. Further, endothelial cells (HUVEC) and neuroblastoma (PC12) cells were seeded in the lumen and surface of the conduit and cultured for 5 days. As the final conduit had higher porosity (80.3 ± 1.4%), it aids in providing a nutrient-nourished microenvironment, promoting neural and vascular ingrowth throughout the conduit. On day 5, the cultured HUVEC and PC12 cells in the triple-layered nerve conduits were found to be viable, proliferative, and uniformly distributed all over the conduit [
Another study by Yoo et al., combined extrusion bioprinting and the electrospinning method to fabricate a longitudinally oriented conduit using collagen and poly(lactide-co-caprolactone) (PLCL) polymers for peripheral nerve defects. A 5% PLCL porous sheet was initially prepared using an electrospinning technique with a uniform thickness and pore size of ∼91 µm and 2.7 ± 0.6 µm, respectively. Two layers of type I collagen bioink were then extruded in a rectangular pattern (8 × 4.7 mm) over the PLCL sheet. The printed collagen patterns were then crosslinked using ammonia vapors. The PLCL/collagen sheet was later rolled and fixed with tissue adhesive to achieve a tubular conduit with dimensions (inner diameter: 1.5-2.0 mm and length:10 mm) equivalent to the rat sciatic nerve. In vivo implantation of PLCL/collagen conduit in an 8 mm long sciatic nerve defect-induced in rats for 12 weeks showed dense and linearly organized axons intraluminally with the expression of neural markers (S-100 and β-tubulin) and thick remyelination. Further, better motor functional recovery observed for the PLCL/collagen-printed group was equivalent to autograft groups, which was determined using ankle contracture angle and tetanic force measurements. Immunohistochemical analysis of the axons within the 12-week implanted conduits also exhibited more regenerated neurons with thicker myelinated axons and significant expression of Schwann cell markers such as S-100 and β-tubulin. Thus, the authors concluded that the NGCs developed using combinational approaches enhanced directed neural growth by the printed collagen, allowed better nutrient diffusion, and outer PLCL electrospun membrane prevented unwanted fibroblast infiltration [
4D printing, an extended version of the additive manufacturing technique, was proposed by Prof. Skylar Tibbits and Prof. Jerry in 2013 with several applications in electronics, robotics, and medical sectors [
]. It is a time-dependent process where the 3D-printed construct transforms into another structure in response to external stimuli such as pH, temperature, voltage, magnetic field, biological factors, etc. Shape-morphing or active origami materials can be 3D printed with programmed anisotropies (e.g., crosslinking density, concentration, alignment of additive components) and allowed to transform their shape/property under the appropriate external stimulus [
]. Wu et al., formulated a hybrid bioink composed of polyurethane, gelatin methacrylate, and gelatin, which exhibited thermo-responsive, photo-responsive, shape memory, and self-healing properties. The prepared polyurethane/gelatin methacrylate/gelatin (PUGG) hydrogel was extruded successfully into various patterns such as mesh, hollow cylinder, sheet, flower, and honeycomb with good printability and shape fidelity. The sheet-like (seven-layer), flower-like (four-layer), and honeycomb-like (four-layer) 3D-printed acellular PUGG constructs displayed the shape recovery property with good structural integrity after immersion in water at 37 °C. In vitro culture of the mouse neural stem cells in the developed PUGG hydrogel showed good viability (94.8%) and proliferation ability (∼3.7-fold increase) compared to 2D cultured cells after 14 days. In addition, the mesenchymal stem cell (MSC)-laden PUGG extrusion-printed 3D constructs were immersed in the cryopreservation agent (7% glycerol/93% cell culture medium) and stored at −20°C or −80°C for three days. In vitro culture of the cryo-preserved PUGG cell-laden constructs demonstrated the shape-recovering behavior at 37 °C with cell proliferation, indicating their shape memory, cytocompatibility, and cryopreservation properties [
]. In another study, a multi-responsive graphene/soybean oil epoxidized acrylate-based (G/SOEA) hydrogel was used to fabricate a self-rolling NGC with high curvature and flexibility via stereolithography technique. The SLA-printed G/SOEA sheet immediately and autonomously wrapped the injured nerve model when placed at 37°C, which confirmed the 4D transformation of the hybrid conduit at the surgery site. In vitro culture of hMSCs on the 4D printed G/SOEA hybrid conduit showed better cellular alignment (due to the printed architecture) with the expression of neural differentiation markers (tubulin-β3, neurofilament heavy polypeptide) and neuron-specific genes (Ngn2, ND1, NSE, and TAU), demonstrating its potential for neurogenic ability [
]. Similar to 4D printing, another recent expansion in AM technique is five-dimensional (5D) printing, which allows the print head and the printed object to move freely at five different angles, developing strong and curved architectures. This sophisticated technology has been used recently for orthopedic applications and can be extended to applications in drug testing and in vivo tissue regeneration. [
3. Design considerations for additive-manufactured (AM) nerve conduits
Peripheral nerve guidance conduits (NGCs) should ideally connect the proximal and distal ends of the injured nerves to promote successful regrowth of nervous tissue with complete functional and motor activity [
] - (i) Material biocompatibility with the host tissues as the materials (such as polymers, cells, growth factors and bioactive motifs) should not have any toxicological and immunological reactions; (ii) appropriate permeability to aid nutrition and gaseous exchange thereby improving the viability of supportive cells (if added) and enhancing vascularization and regeneration process; (iii) adequate physical and mechanical properties with native-equivalent flexibility to avoid plausible collapse, compression or breakage of the conduit under physiological environment; (iv) biomimetic intraluminal structures to provide physical guidance specific for motor and sensory axonal regeneration; (v) cell-supportive environment for effective signal transmission; (vi) appropriate degradation ability as the nerve conduit must remain intact during all stages of nerve regeneration and then degrade gradually with negligible non-toxic byproducts, which is required to avoid secondary surgery and (vii) suturable, sterilizable and transparent (desirable by clinicians, however not a necessary property) with ease of handleability. Until now, several research reports have shown the successful fabrication of PNCs using conventional scaffold fabrication approaches and proved nerve ingrowth using the developed conduits in several in vivo animal models[
], yet, these PNCs have not been commercialized, which may be mainly due to the use of complex laboratory-scale manufacturing process, time-consuming technologies and suboptimal efficacy in matching the structural, biochemical and mechanical property with the host tissue. Hence, a better understanding and appropriate interplay between the ideal NGC properties are necessary to recapitulate the microarchitectural, structural, and functional complexity of a native nerve. Thus, this section presents a detailed overview of several design criteria (such as choice of bioinks, printing parameters, porosity, degradation rate, mechanical strength, and design architecture) that are involved in the fabrication of ideal NGC using additive manufacturing techniques to foster successful peripheral nerve regeneration.
3.1 Choice of 3D printable materials
The choice of a biomaterial is an important criterion as it greatly impacts the 3D printed nerve conduit properties, including chemical resistance, printability, permeability, degradation, mechanical strength, flexibility, and sterilization stability [
]. Generally, the 3D printed / bioprinted PNC are fabricated using various polymers, including natural polymers, synthetic polymers, and hybrid (natural/synthetic) polymers as printable ink or biomaterial inks. These polymers are used as hydrogels, polymeric filaments/pellets, or low-viscous polymeric solutions for constructing 3D constructs. Incorporating or functionalizing these polymers or the printed 3D constructs with bioactive peptides (e.g., IKVAV, RGD), supportive cells (e.g., Schwann cells, DRGs), growth factors (e.g., nerve growth factor, neurotrophins, BDNF) and conductive components (e.g., graphite, carbon nanotubes (CNTs), PPy) have accelerated the regeneration of peripheral nerves. Many researchers have shown interest in developing PNCs using medical-grade polymers as these polymers are considered biocompatible and biodegradable with low-toxic byproducts [
]. However, there are no regulatory definitions and perspective requirements for "medical-grade polymers and their devices," specifically for raw material synthesis, manufacturing protocols, and commercializing medical devices using these polymers [
]. In addition, physical or chemical modifications of these materials would enhance or tune the material properties similar to target tissues such as nerves, cardiac, and muscles, which requires scrutinization during the material selection and scaffold fabrication process [
]. Further, the choice of polymers depends upon the manufacturing process as they require unique bioink properties for fabricating 3D constructs in minimum printing duration and post-processing steps, which are crucial determinants for scalability. For instance, the extrusion-based and laser-assisted bioprinting methods use high/low viscous polymeric solutions as bioinks with properties including viscoelasticity, shear thinning ability and instantaneously cross-linkable or in-situ gel-forming ability. On the other hand, the bioinks / biomaterial inks used for printing via stereolithography should possess low viscosity and photo-crosslinking ability when irradiated at specific conditions to obtain stable 3D structures.
3.1.1 Natural polymers
Natural polymers are derived from renewable resources such as plants, animals and microorganisms and have been used for fabricating 3D tissue scaffolds. The most commonly used natural-based polymers include alginate, gelatin, gellan gum, hyaluronic acid, silk fibroin, and collagen, as they possess significant advantages such as good biocompatibility, tunable degradation kinetics, chemically tailorable to introduce cell-binding motifs (if required), anti-immunogenic and bioinert nature [
]. In addition, polymers, such as collagen, gelatin, and dECM extract from any desired tissues, may resemble native ECM architecture and chemical backbone. However, they have poor mechanical properties, which may collapse the structural stability of the printed constructs when used for long-term implantation. Hence, two or more natural polymers are blended or modified using physical, enzymatic, or chemical methods to obtain a stable NGC. For example, Wu et al., bioprinted a photo-crosslinkable natural polymer-based nerve construct using an extrusion-based 3D bioprinter. Optimized concentrations of biopolymers such as gelatin methacrylate (GelMA), alginate (Alg), and bacteria nanocellulose (BNC) were blended initially and mixed with 0.5% of photoinitiator lithium phenyl(2,4,6-trimethyl benzoyl) phosphinate (LAP) to achieve properties such as shear thinning ability, ease of crosslinking and biocompatibility. The blended polymers were then printed into a cuboid of 8 × 8 × 2 mm and a 5 × 4 mm cylinder structure. Later, the printed structures were photo-crosslinked using blue light at 430 nm for 20 s, followed by physical crosslinking with a sterile 50 mM CaCl2 solution. The printed structures showed clear, sharp-edged printed patterns at optimized printing conditions and maintained their shape throughout the study. The conductivity of the blended polymer (Alg/GelMA) increased with the addition of BNC (Alg/GelMA/BNC) due to the presence of hydroxyl groups in BNC and the formation of nanopores within the formed hydrogel, which are necessary for an ideal nerve conduit. In addition, the stress-strain curve obtained for the developed hydrogels was similar to that of the native soft tissues. The pre-hydrogel solutions made up of alginate, GelMA, and BNC were printed along with rat Schwann cells (RSC 96), followed by physical and photo-crosslinking, and cultured in vitro for 7 days. The cell-laden printed constructs showed increased cell viability, proliferation and promoted neuronal expression (such as S100β, cytoskeletal staining) with spindle morphology at day 4. The gene expression studies using quantitative real-time PCR (qRT-PCR) and PCR array revealed an increase in the expression of genes responsible for neuronal proliferation and myelin formation, such as ASCL1, POU3F3, NEUROG1, DLL1, NOTCH1, and ERBB2 at day 7 in the Schwann cell-laden Alg/GelMA/BNC constructs compared to the bioprinted constructs without BNC. Subcutaneous implantation of RSC96 cell-laden Alg/GelMA/BNC printed constructs in nude mice for 4 weeks also exhibited expression of S100β, confirmed using immunohistochemical staining. Thus, alginate/GelMA/BNC may be a suitable natural biomaterial component for developing PNCs, however, poor in vivo degradation of the bacterial cellulose should be considered due to the lack of cellulase enzymes in the native environment [
Similarly, collagen, a naturally occurring polymer, is one of the body tissue's most abundant extracellular matrices. Due to the limitations such as faster degradation properties, poor mechanical stability, and high cost, collagen is often blended with other natural polymers to improve the required properties. Li et al., blended collagen with silk fibroin (isolated from silk worms) in an equal mass ratio and extruded using the following printing parameters - printing speed of 9 mm/s, solution extrusion speed of 2 mm/min, and layer thickness of 0.1 mm. The freeze-dried printed 3D scaffolds displayed microporous and honeycomb-like structures. These constructs were implanted in the cystic space of female SD rats to study the degradation behavior for up to 4 weeks. The implanted scaffolds showed 20%, 59%, and 74% degradation after 1, 2, and 3 weeks, respectively. At week 4, the complete integration of the scaffold with surrounding tissue was observed, which showed that the scaffold had regenerated the tissue successfully. Further, an injury model was developed by removing a 3 mm cord at the T10 region of the spinal cord. The 3D-printed scaffolds were then implanted immediately at the lesion gap to evaluate the regeneration potential of the 3D-printed collagen / SF (3D-C/SF) scaffold. The locomotory recovery test (Basso, Beattie, and Bresnahan (BBB) open-field test) showed a BBB scoring of 7.91 ± 0.68 with higher amplitude and shorter latency for 3D-C/SF implanted group at 8 weeks after surgery. The MRI scan of the 3D-C/SF implanted group showed partial reconnection between the proximal and distal sites, which confirmed the recovery of the spinal cord injury. In addition, the 3D-C/SF scaffold implanted group showed positive for neuronal factor expression in nerve fibers wrapped by myelin sheath after eight weeks. Overall, the scaffold supported the regeneration of nerve fibers after spinal cord injury in a shorter implantation period [
] were also used for fabricating peripheral nerve conduits to guide the neural cells and promote nerve regeneration. Unlike natural polymers, synthetic polymers show increased mechanical properties, ease of functionalization, and structural integrity [
]. Yet, they lack cell-adhesive properties, which can be improved by functionalizing or blending cell-binding motifs/drugs to enable cell adhesion and proliferation. More particularly, PU-based conduits can be fabricated easily with tunable mechanical properties with structural integrity, having better cellular effects in terms of viability, proliferation, and protein expression than traditionally fabricated PU scaffolds [
]. Chen et al., have recently fabricated 3D-printed nerve conduits using Astragaloside (Ast) containing water-based PU resin via a UV-based DLP 3D printer. Initially, acrylate-functionalized and water-based PU was prepared by reacting PU resin with 1.5% 2,4,6-trimethyl benzoyl-diphenyl-phosphine oxide, 0.1% 2-hydroxy-4-methoxy benzophenone-5-sulfonic acid hydrate, 0.01% 4-hydroxy-2,2,6,6-tetramethyl piperidinooxy, and 30% 2-hydroxyethyl methacrylate. Various concentrations (10 and 20 µM) of Ast, a Chinese drug aid in Schwann cell regeneration, were added to the synthesized photoactive PU resin (Ast/PU) and fabricated the hollow nerve conduits as per the 3D design (14 mm length, 2 mm diameter). Increasing the concentration of Ast in PU resin increased the hydrophilicity (contact angle of 20 µM Ast/PU – 68.6±2.3°) and degradability (of 20 µM Ast/PU ∼21% at 4th week) of the 3D-printed synthetic nerve conduits than non-drug PU conduits and 10 µM Ast/PU conduits. The primary rat Schwann cells (RSCs) were seeded on the 3D-printed PU conduits and evaluated the cell morphology and proliferation at various time points (1, 3, and 7 days). Ast/PU conduits showed a significant increase in cell proliferation compared to plain PU conduits. In addition, Ast-containing PU conduits displayed cells with longer axons and spindle-shaped morphology on day 7. Further, the 20 µM Ast/PU conduits showed increased expression of BDNF, HuC/HuD proteins, NGF, and decreased HMg-Box gene-10 (SOX10) expression compared to other groups, such as 0 µM Ast/PU and 10 µM Ast/PU. Overall, the 3D-printed 20 µM Ast/PU nerve conduits showed good mechanical properties, enhanced neural cell proliferation and regeneration, and increased expression of neural-related markers, which may be suitable for peripheral nerve-regeneration-related applications [
Researchers have attempted to develop hybrid polymers by physically blending or chemically functionalizing both natural and synthetic polymers, thereby enhancing the advantages of both natural and synthetic polymers and reducing their limitations. It is known that synthetic polymers lack cell adhesion properties and show poor hydrophilicity, decreased biodegradability, and higher mechanical properties. Hence, by blending or conjugating natural polymers, the limitations of synthetic polymers may be avoided, resulting in desirable properties for the successful fabrication of biological constructs. For example, Zhu et al., have blended natural (GelMA) polymer and synthetic (PEGDA) polymers with photoinitiator (LAP) dissolved in Dulbecco's phosphate-buffered saline and developed a peripheral nerve guidance 3D construct using DLP-based 3D printing technique. The authors have fabricated linear, tubular with single- and four-channel conduits and branched conduits using optimized concentrations of GELMA (7.5%), PEGDA (25%), and LAP (1%). GelMA-based hydrogels, though possessing cell-binding motifs and photo-crosslinkable features, are considered soft and fragile, requiring blending with other polymers (e.g., PEGDA) to obtain desired mechanical properties. In addition, by varying the concentration of LAP (0.2 and 1%), degree of methacrylation in GelMA, and crosslinking light intensity (6.7 and 16.6 mW/cm2), the mechanical property of the fabricated GelMA/PEGDA 3D constructs can be tuned to achieve Young's modulus (4.5 MPa @ 1% LAP and 16.6 mV/cm2) equivalent to the nerve tissues tissue (0.5–13 MPa). In vivo implantation of the four-channel conduit in a 6 mm sciatic nerve defect in transgenic mice demonstrated directed ingrowth of regenerated nerves from proximal to the distal end of the 3D construct with faster recovery of motor function, sensation and neurofilament expression within 11 weeks of surgery [
Similarly, Liu et al., utilized a multi-nozzle-based pneumatic extruder to develop a bilayered nerve conduit by spirally depositing GelMA and poly(ethylene glycol) diacrylate (PEG-DA) polymers over the rotating rod. The inner and outer layers of the 3D-printed conduits were primarily composed of 5% (w/v) GelMA mixed with bone marrow mesenchymal stem cells (BMSCs) and GelMA/PEGDA polymer along with the photoinitiator (0.5% LAP), respectively. The inner cell-laden layer was photocrosslinked under UV light at low intensity (0.5 W/cm2) to obtain soft hydrogels allowing BMSCs to proliferate extensively with an elongated morphology. In vitro seeding of PC12 cells inside the lumen of the printed 3D conduit also enhanced cell viability with spindle-shaped proliferating cells at day 3 itself. In contrast, the outer layer (GelMA/PEGDA) was UV-crosslinked at higher light intensity (2 W/cm2), which resulted in increased mechanical strength (Young's modulus ∼2.9 MPa; Compressive modulus ∼3.4 MPa) than plain GelMA hydrogel layer (Young's modulus – 318.46 ± 29.26 kPa; Compressive modulus – 304.15 ± 44.88 kPa). The outer stiff printed layer also enabled 3D structures with good shape fidelity and structural integrity, thereby preventing the collapse of the conduits [
]. Several reports emphasize the flexibility in using hybrid (natural-synthetic) polymers to develop native-equivalent NGCs with tunable processability and printability that eventually accelerate axon regeneration in the nerve defect site.
3.1.4 Additive components
Biological factors such as live cells, peptides, drugs, growth factors and cytokines were encapsulated in nanocarriers or added directly to the polymeric inks to regulate the attachment, proliferation, migration and homeostatic pathways of the neurons in the injured site, thereby accelerating the tissue regrowth and recovery within the conduit. These biological additives in the biomaterial inks also influence the rheological and printable properties, such as increasing the solution's viscosity and the printing forces (e.g., pressure, laser energy) during the deposition process. Tao et al., have fabricated a GelMA-based 3D nerve conduit loaded with hippo pathway inhibitor encapsulated in methylated poly (ethylene glycol)-poly(3-caprolactone) (MPEG-PCL) nanoparticles. Initially, a hippo pathway inhibitor called XMU-MP-1 (4-((5,10-dimethyl-6-oxo-6,10-dihydro-5H pyrimido[5,4-b] thieno[3,2-e] [1,4]diazepin-2-yl) amino) benzene sulfonamide) and core-shell MPEG-PCL nanoparticles were synthesized separately, and the drug XMU-MP-1 was encapsulated into the MPEG-PCL nanoparticles. Later, these drug-loaded NPs (100 µg/mL) was mixed with 20% gelatin-methacryloyl (GelMA)/1% LAP pre-hydrogel solution. In this study, the synthesized hydrophobic drug inhibitor XMU-MP-1 aid in downregulating the hippo pathway and upregulating the yes-associated-protein (YAP) expression, thereby controlling myelin elongation. The DLP 3D printer was used to fabricate a 13 mm hollow nerve conduit using the prepared bioink, which was further crosslinked under UV exposure (405 nm) for 90 s. The incorporated drug XMU-MP-1 was mixed at various concentrations (0.1–0.5 mg/mL) to the pre-gel solution, where 0.5 mg/mL of the drug promoted the faster migration and proliferation rate (11%) of Schwann cells compared to other concentrations, which was confirmed using in vitro wound healing assay. Direct treatment of S16 cells with 0.5 mg/mL of XMU-MP-1 drug has predominantly increased the gene expression of neuronal markers such as BDNF, NGF, neurotrophins-3 (NT-3), fibroblast growth factor-2 (FGF-2), connective tissue growth factor (CTGF) and cysteine-rich angiogenic inducer 61 (CYR61) compared to low drug concentration groups. Further, the drug-release kinetics from the fabricated GelMA conduit showed a reduced drug release (47.4%) in contrast to the pattern observed for pristine drug-encapsulated NPs, which showed higher release (∼83.2 %) after 144 h. Subcutaneous implantation of the 3D-printed conduits in SD rats was stable for 12 weeks, which may be sufficient for successful neural regrowth. Moreover, the electrophysiological analysis of the implanted drug-loaded conduits in a 10 mm rat sciatic nerve defect for 3 months demonstrated higher nerve conduction velocity (NCV), latency, and peak amplitude of compound muscle action potential (CMAP) compared to non-drug-loaded conduits. In addition, morphological analysis of the regenerated nerves also showed axon diameter and myelin sheath thickness similar to autograft groups [
In another study by Tao et al., live platelets incorporated hydrogel-based nerve conduits had significantly promoted native axonal growth due to the sustained release of multiple nerve-regenerating growth factors such as tumor growth factor-β, platelet-derived growth factor, fibroblast growth factor, and vascular endothelial growth factor by the activated platelets. In this study, DLP-based 3D nerve conduits were prepared using gelatin methacrylate/poly (ethylene glycol) diacrylate (GelMA/PEGDA) bioink containing live platelets at various concentrations (6.25–100 × 106 cells/mL). The encapsulated platelets were stimulated by adding thrombin in the bioink solution to enhance the release of growth factors. In addition, the inactivated platelets showed prolonged survival and sustained release of growth factors when activated. In vitro culture of Schwann cells in the GelMA/PEGDA hollow conduit aided in better proliferation and upregulation of TGF-β signaling genes such as Crlf1, Fstl3, and Tagln compared to the non-platelet loaded conduit. Further, in vivo implanted platelet-encapsulated 3D conduits in rats exhibited a higher density of myelinated nerve fibers with increased myelin sheath thickness, axon diameter, and GAP-43 expression, which were equivalent to the autograft group, thus confirming its regenerative ability in peripheral nerve injury [
In addition to the biological factors, components providing conductivity, such as graphene oxide, CNTs, PPy, polyaniline, polythiophene, and poly(3,4-ethylene dioxythiophene), were incorporated with polymeric biomaterial inks to induce conductivity and accelerate nerve regeneration. For example, Uz et al., have developed a gelatin-based 3D bioprinted nerve conduit containing graphene as an electrical conducting agent, which possesses ECM-mimic microstructures and mechanical strength that aid in the transdifferentiation of MSCs into Schwann cells. A 3 mm diameter graphene rod integrated gelatin-based nerve conduits were prepared initially using conventional methods such as molding and thermally induced phase separation technique, which was then chemically crosslinked via 1-ethyl-3-(3-dimethyl-aminopropyl) carbodiimide-N-hydroxy-succinimide (EDC-NHS) chemistry. Additionally, an FDM-based graphene/PLA-based interdigitated circuit was 3D-printed and embedded in a gelatin scaffold to develop a 3D-printed graphene-based gelatin conduit. The conventional and 3D printed conduits displayed similar microstructure with 90% porosity with an average pore size of ∼135 μm and a swelling ratio of 830%. The mechanical strength of the conduit was evaluated in the absence of graphene rods and circuits, and the results showed that pristine gelatin scaffold exhibited elastic properties bearing Young's modulus, yield strength, and average complex modulus of ∼80 MPa, ∼1.8 MPa, and 7.6 × 106 Pa, respectively. Further, the graphene rod/circuit-integrated gelatin nerve conduit offered a resistance of up to 50 kΩ, sufficient to glow the LED bulb, thereby confirming its conducting ability. In vitro seeding and culture of MSCs at a density of 5 × 105 cells/mL on the gelatin-based scaffold for 15 days demonstrated good attachment and proliferation within the gelatin construct. In addition, applying electrical stimulation upto 100 mV at 50 Hz for 10 minutes continuously for 10 days allowed transdifferentiation of seeded MSCs over the gelatin-based conduits integrated with graphene rods and 3D-printed circuit into Schwann cell-like phenotype, which was confirmed by the expression of functional neuronal markers such as S100, S-100β, and P75 [
The structural design/cues in the developed PNCs are also an important parameter for an ideal nerve conduit as it physically supports and directs the growth between the host and newly regenerated axons in the injured site. The simplest structure used in PNCs is the single-lumen/hollow conduits or wraps with desired dimensions, prepared using conventional scaffold fabrication techniques. However, they fail to direct the growing axons towards the distal end and limit the exchange of nutrients throughout the conduit. Hence, these conduits were luminally filled with structural, hormonal, or cellular components (e.g., fibrin, collagen, laminin, neurotrophic factors, autologous/syngeneic neural cells), providing a larger surface area for cell growth and proliferation, mimicking fascicular architecture and overcoming diffusion limitations. Yet, these lumen-filled conduits may not be clinically applicable due to few limitations, such as the controlled release of embedded factors, material availability, and batch-to-batch variability during the preparation process [
]. Despite these limitations, NeuraGen® 3D nerve guide matrix is the only commercial nerve conduit with an outer collagen membrane filled with porous collagen matrix infused with chondroitin-6-sulfate. However, there are no published reports for both pre-clinical and clinical results for NeuraGen® 3D nerve guide matrix product. AM techniques demonstrate great potential in developing personalized PNCs with controlled porous architecture and uniform surface area, enabling directed and accelerated neural regeneration. The common design strategies for developing 3D-printed conduits include unbranched single-channel conduits, multi-channel conduits, branched conduits, irregularly shaped structures (e.g., grids), and anatomically nerve-equivalent 3D structures [
]. These 3D models may partially mimic the native nerve dimensions, architectures, and physiological properties, including mechanical strength, thus confirming its superiority over conduits prepared using traditional fabrication methods.
3.2.1 Unbranched single-channel conduits
The simplest design for fabricating 3D NGCs is the linear non-porous hollow conduits with personalized dimensions, which do not require complex geometries and can be developed in low production time on an industrial scale. A few studies have suggested that these conduits, though printed with nerve-equivalent dimensions, were disadvantageous due to improper nutrient exchange, poor mechanical strength, poor cell attachment with misguided growth, incomplete nerve regeneration, and unwanted tissue ingrowth, more particularly observed in longer nerve gap defects. In addition, hollow conduits were only beneficial for smaller and shorter nerve gaps (<30 mm) with non-satisfying functional recovery in animal models [
] (such as collagen, hydrogels, cells, drugs, and neurotrophic growth factors) in these conduits, thereby proving their potential towards enhanced axonal alignment and regrowth. For instance, Zhang et al., developed a DLP-printed self-adhesive wrap using photocrosslinkable monomers such as azide-modified gelatin methacrylate (N3-GelMA) and drug (XMU-MP-1) loaded dibenzyl cyclooctyne-modified gelatin methacrylate (DBCO-GelMA). These monomers were 3D-printed in rectangular patterns and allowed to self-adhere the printed layers, followed by a rolling process to form a wrap-like hollow conduit (Fig. 8A). The incorporated drug (XMU-MP-1) showed sustained release in vitro and can potentially influence the growth and migration of Schwann cells. Further, in vivo implantation of 3D-printed nerve wraps in sciatic nerve transected rats also showed faster nerve conduction velocity (NCV – 36.7 ± 3.8 m/s) and shorter latency time of compound motor action potential (CAMP – 1.93 ± 0.20 s) than the end-to-end neurorrhaphy group (NCV – 27.9 ± 3.4 m/s; CAMP – 2.55 ± 0.29 s), indicating better regeneration and functional recovery in the drug-loaded nerve wraps [
The main objective of developing ideal NGCs is personalized nerve tissue architecture incorporated with adequate vasculature and multiple biological components to promote faster axonal growth and maturation. With the advancement in machine intelligence and 3D imaging techniques with computational design and analysis, several NGC architectures were designed and analyzed using 3D modeling and analysis software. Thus, the physical and mechanical properties of the designed NGC structures can be predicted and easily modified as per the chosen nerve properties, which would ultimately reduce the bottlenecks in the manufacturing process [
]. Hence, Zhang et al., adopted finite element computational analysis to identify an ideal unbranched NGC structure from different NGC conduit designs – hollow luminal (HL NGC) with circular and square pores, micro-grooved (MG NGC), and multichannel (MC NGC). The authors stated that the computational analysis of the modeled 3D conduits might provide a customizable solution to reinnervate nerves with optimized and desired properties (such as porosity, surface area/volume, moduli values, and permeability) essential for an ideal NGC. It is known that the pore size and porosity directly influence the mechanical properties and permeability, which are also important criteria for designing an ideal NGC. The computational results suggested that HL NGCs with uniform circular pores had better mechanical strength (tensile modulus – 8.1 MPa, bending modulus – 2.3 MPa) but were more poorly permeable (porosity – 59.8% and permeability – 3.27 × 10−9 m2) than squared pores (tensile modulus – 8 MPa, bending modulus – 0.9 MPa, permeability – 3.39 × 10−9 m2 and porosity – 64.2%). In addition, the pores in the longitudinal direction of NGCs may biologically favor the cell alignment more than the circumferential pores of NGCs. Further, MG NGCs and MC NGCs had similar tensile modulus values, however, MC NGCs were impermeable (0 × 10−9 m2) compared to MG NGCs (2.63 × 10−9 m2). Among three different designs, MC NGCs were observed as disadvantageous since they display zero permeability value though having higher stiffness values and larger surface area, aiding cell attachment [
Another strategy for improving the neural regeneration process is the fabrication of NGCs with unbranched multi-channels with uniform and longitudinal channel pores throughout the conduit. The intraluminal dimensions of multi-channels in the conduit resemble the individual nerve fascicles, guiding the alignment and growth of the Schwann cells and axons. These conduits were developed predominantly using SLA[
] techniques, which suggest the positive effects of 3D-printed multi-channeled nerve conduits with uniform-sized (e.g., diameter) channels via in vitro and in vivo neural regeneration studies, where the newly formed axons were present throughout the conduit with a thick myelin sheath formation [
]. Krieghoff et al., produced extrusion-based multi-channeled NGCs with rectangular and circular peripheral structures using biocompatible gelatin polymer and anhydride functionalized oligomers (oligo(pentaerythritol diacrylate monostearate-co-maleic anhydride-co-N-isopropyl acrylamide) and oligo(pentaerythritol diacrylate monostearate-co-maleic anhydride-co-diacetone acrylamide)) (Fig. 8B-ii). Initially, the gelatin hydrogel premixture solution was prepared by mixing gelatin and anhydride-containing oligomers and loaded into the extrusion syringe. The gelatin pre-gel mixture was 3D-printed at a speed of 6 mm/s and 8 mm/s for rectangular (10 × 11 mm, 50 layers) and circular (13 mm diameter, 50 layers) outlines, respectively, exhibiting different shaped outlines with porous features. The printed structures were instantaneously crosslinked (click-chemistry reaction) between the anhydride groups of oligomers and amine groups of gelatin under the presence of an inorganic catalyst, potassium hydrogen phosphate (K2HPO4) to achieve scalable and reproducible multi-channeled conduits. The printed hydrogel-based conduits showed slow hydrolytic degradation behavior with a good proliferation of human adipose-derived stem cells with an elongated morphology. The authors stated that these conduits may be suitable for peripheral nerve conduits as they provide a large surface area for cell attachment and growth [
]. Further, studies on the fabrication of varied lumen diameters and shapes of multi-channels in these conduits may be appropriate as they resemble exactly the fascicular structures present in the native nerves. However, limitations in the diffusion properties in the multi-channeled conduits restrict their use clinically though having higher mechanical strength than single-lumen conduits.
3.2.3 Branched conduits
Peripheral nerves are mostly branched with interconnected networks and varied dimensions (mostly tapered towards the end), requiring the development of similar tubular structures, thereby limiting the harvest of autologous nerve grafts. Generally, the anatomically nerve-equivalent conduits can be ideally fabricated using the following crucial steps – (i) identification of the damaged nerves (target nerve) using high-resolution imaging techniques; (ii) generation of a 3D model corresponding to the identified damaged nerve and (iii) fabrication of branched NGCs using appropriate 3D printing parameters associated with the chosen AM technique. Johnson et al., have followed similar steps to develop an anatomically-equivalent bifurcated conduit by combining 3D imaging and 3D printing techniques (Fig. 8C). The bifurcated sciatic nerve tissue with sural and tibial branches was initially identified, transected, and imaged intact using structured light scanning (SLS), resulting in a 3D nerve geometrical model. The reconstructed 3D model was then printed layer-by-layer using silicon material with axially-oriented grooves in the luminal surface to provide physical cues resembling bands of Büngner to aid Schwann cell alignment. In addition, the bifurcated nerve model was printed path-specifically using gelMA/NGF and gelMA/GDNF, allowing sensory and motor cues in a concentration gradient manner, respectively. Further, these path-specific 3D conduits were successfully implanted in a sciatic nerve defect in rats to observe the regenerative potential of the conduit after three months. The sciatic nerve was sutured to the proximal end of the conduit, and the tibial and sural nerves were sutured with the conduit's distal motor and sensory cue channels. After 3 months of implantation, histological examination displayed increased nerve growth with S100 expression within the conduits. Gait analysis of rats also showed significant improvement in the rat walking behavior with a gait duty cycle of factor ∼1.4. Thus, this study demonstrated the feasibility of fabricating personalized 3D-printed / bioprinted constructs with complex branches, dimensions, fascicular features, mechanical and biological cues, possibly favored using high-resolution 3D scanning/imaging and printing techniques [
]. Similarly, Zhang et al., utilized the DLP 3D printing method to develop multi-branched tubular networks extended from a single nerve stump using 20% gelMA as biomaterial ink, 0.15% vitamin B12 as photo absorber to absorb excess light and 0.7% lithium phenyl-2,4,6-trimethyl-benzoyl phosphinate (LAP) as a photoinitiator. The printed 3D structures were photo-crosslinked at 405 nm to obtain mechanically stable multi-branched networks with good printing fidelity (Fig. 8C) [
]. Another study by Ramesh et al., fabricated PLA-based 3D nerve constructs using the FDM technique by mimicking the fascicular bundles obtained from a magnified region of interest (ROI) of micro-CT scanned peripheral nerve. The 3D-printed constructs mirrored the architecture of the neural layers (endoneurium and perineurium) and nerve fascicles, confirmed using SEM images. Incorporating bovine serum albumin protein-based nanoflowers over the printed constructs aided the seeded PC12 cells with neurite extensions and higher expression of neural-specific markers (Neurofilament-200 and beta-tubulin). However, these constructs were non-flexible, requiring in vivo implantation with a detailed investigation of biocompatibility and motor recovery studies [
In addition to tubular and branched constructs, flat-patterned structures such as grids with uniform porous architecture have also been 3D-printed to evaluate the technique's feasibility in printing the bioinks / biomaterial inks, physicochemical and biological properties of the printed scaffold as an initial proof-of-concept, which can later be altered to develop 3D-printed structures in any desired shapes. Techniques such as stereolithography[
] were most commonly adopted for developing mesh-like 3D scaffolds and later rolled or printed to form tubular constructs. In a study by Wu et al., the authors have fabricated an extrusion-based grid-patterned scaffold using alginate/gelatin/rat Schwann cells with optimized printing parameters such as an extrusion speed of 0.15 mL/s, a nozzle scanning speed of 3.5 mm/s, nozzle diameter of 160 µm and printing temperature of 28°C. The printed constructs were then ionically crosslinked using 50 mM CaCl2 for about 5 minutes. In vitro culture of the alginate/gelatin/RSC constructs displayed good viability and proliferation for about 7 days, with the printed Schwann cells adhering well to the scaffolds with elongated neurites. The gene expression of neurotrophic factors such as BDNF, NGF, platelet-derived growth factor (PDGF), and glial-derived neurotrophic factor (GDNF) in the cell-laden printed 3D scaffolds was higher compared to the cells cultured in the 2D environment. Further, subcutaneous implantation of the bioprinted grid-shaped scaffold in rats for four weeks revealed the presence of few inflammatory cells and showed a positive expression of the S-100β marker. Later, the alginate/gelatin biomaterial ink was used to print cylindrical channeled conduits with 5 mm height and 4 mm diameter, enabling the development of tubular 3D constructs (Fig. 8D). However, in vitro and in vivo studies were not performed using tubular constructs, which may be due to the poor stability and swelling nature of alginate/gelatin biomaterial ink [
Permeability is another parameter that regulates the diffusion of proteins, essential nutrients, neurotrophic / growth factors, and gaseous exchange for faster regeneration with vascularization in deeper areas of the conduit. The optimum porosity required for a peripheral nerve conduit is about 60-80%, with a pore size ranging between 4-30 µm, which is sufficient enough to prevent the infiltration of non-neural cells such as fibroblasts within the conduit and enhance the viability of supportive cells such as Schwann cells [
]. The pores in the NGCs should also primarily support initial fibrous cable formation and enhance cell regulatory activities during the regeneration process. However, the porosity of the conduits affects their mechanical properties, such as flexibility, stiffness, and resistance to kink and collapse, which are very important during in vivo implantation and under physiological stresses. Most 3D-printed and bioprinted PNCs were observed to be semipermeable with interconnected pores, thereby promoting neuronal regeneration within the conduit. Lee et al., have developed photo-crosslinkable bioinks composed of 40% polyethylene glycol (PEG), 60% polyethylene glycol diacrylate (PEG-DA), 0.5% photoinitiator (Irgacure 819), and incorporated various concentrations (0.1%, 0.5%, and 1.0%) of PLGA nanoparticles. Using the SLA technique, the highly aligned 3D scaffolds were developed using the prepared bioink with different geometrical porosities such as small (44%), medium (56%), and large square (68%) per the CAD design. Similar porosities with uniform interconnected porous structures were observed from SEM images of the SLA-printed constructs. PC12 cells were seeded at a density of 30,000 cells/scaffold and allowed to culture for 4 h to evaluate the cell attachment over the scaffolds. Results suggested that only large porous scaffolds (68%) showed significantly higher cell attachment due to highly desirable nanofeatures for nerve scaffolds which mimic native ECM nanostructures, compared to small and medium porous scaffolds. Further, adding PLGA nanoparticles in the 3D construct improved Young's modulus and ultimate tensile strength by 250% over the 3D scaffold without PLGA nanoparticles. In addition, the culture of primary rat cortical neurons over these scaffolds also demonstrated a higher proliferation rate (upto 7 days) with increased neurite extensions due to the support of micro-grooved topographies, confirmed by the expression of early neural markers such as TUJ-1 and MAP2 [
]. Similarly, the same group had functionalized multi-walled CNTs (MWCNTs) in the photo-crosslinkable PEG/PEGDA/Irgacure ink to fabricate SLA-based electroconductive scaffolds with different porosities (large 66%, medium 52%, and small 31%). Medium porous scaffolds showed consistent and better printing quality than other printed scaffolds (small and large porosities). In addition, in vitro culture of neural stem cells (NE-GFP-4C) in different porous scaffolds under biphasic pulse stimulation for 4 h demonstrated better cell attachment and expression of nerve differentiation markers (Nestin, MAP 2, TUJ1, and GFAP) in the medium porous 3D constructs than other scaffolds [
]. Though not many published reports emphasize the pore organization (uniform/multi-pore) and total porosity within the 3D-printed NGCs, achieving optimum porosity values without affecting the mechanical and biological properties (e.g. fibroblast infiltration) is important for an ideal 3D-printed construct.
Conductivity is another ideal property of nerve, cardiac, and muscle tissues. Developing conducting 3D scaffolds for nerve tissue engineering aid in quickening the regeneration and maturation process. These scaffolds are generally fabricated by introducing conductive polymers such as PPy, reduced/nanographene oxide (rGO/nGO) polyaniline (PANI), CNTs, poly(3,4-ethylene dioxythiophene) (PEDOT), and polythiophene (PTh) at appropriate concentrations for making it biodegradable and biocompatible. Incorporating these materials into the 3D-printed NGCs is feasible for external electrical stimulation, which could enhance cell proliferation, migration, and differentiation of neural cells. In a study by Namhongsa et al., PPy-coated poly(l-lactide-co-ε-caprolactone) (PLCL), and poly(l-lactide-co-glycolide) (PLGA) scaffolds were developed by combining 3D-printing (3D) and electrospinning (E) technique, represented as PLCL-3D/E/PPy and PLGA-3D/E/PPy respectively. Various concentrations of PPy (3.7, 7.4, 15, and 22 mM) were uniformly coated via sonication over the PLCL-3D/E and PLGA-3D/E scaffolds to measure the conductivity property. As PPy concentration increases, the conductivity of the scaffolds was observed to be unstable in addition to non-uniform coating, which may be due to the aggregation of PPy molecules over the electrospun PLCL and PLGA scaffolds. Hence, with the optimized concentration of PPy (7.4 mM), the scaffolds' conductivity was 12.04 ± 0.12 and 10.50 ± 0.08 S/cm for PLCL-3D/E/PPy and PLGA-3D/E/PPy scaffolds, respectively. In vitro culture of Schwann cells on the PPy-coated PLCL/PLGA scaffolds showed good viability (70%) and proliferation with elongated morphology on day 7 compared to non-PPy-coated scaffolds. Moreover, a higher concentration of PPy (aggregated PPy) hindered the cellular infiltration and migration within the scaffold, reflecting poor regeneration. Thus, the authors concluded that the developed scaffold exhibited conductive, biocompatible, and ECM-mimetic features necessary for an ideal NGC [
] carbon hybridized bonds resembling neural topography and enhance neural regeneration when added at suitable concentrations. Lee et al., incorporated the amine-functionalized multi-walled CNT (MWCNT) in PEGDA polymer to develop MWCNT-PEGDA 3D electroconductive grid-patterned scaffold using a stereolithography 3D printer. The developed PEGDA and PEGDA-MWCNT hydrogels showed uniform porous structures, confirmed using SEM images. Various concentrations of MWCNTs (0.02, 0.05, and 0.1%) were blended with PEGDA polymer hydrogels to evaluate the electrochemical properties of the prepared PEGDA-MWCNT hydrogels using cyclic voltammetry. Before the analysis of redox measurements, these hydrogels were immersed in PBS for 7 days to leach out the loosely attached MWCNTs from the hydrogels. Results demonstrated that only 0.1% MWCNT containing PEGDA scaffolds had a significantly higher charge storage capacity of 2.21±0.12 mC/cm2 than the plain PEGDA scaffolds (0.133±0.09 mC/cm2), which may help in providing prolonged conductivity of the scaffolds. In addition, 0.1% MWCNT-PEGDA SLA-based 3D scaffold had a uniform distribution of MWCNT with increased hydrophilicity and Young's modulus (189%) than PEGDA 3D-printed scaffold. In vitro culture of neural stem cells (NSC) on 0.1% MWCNT-PEGDA 3D scaffold showed more than 70% NSC adhesion on day 7 with neurite extension (average length ∼28%), ensuring the scaffold's biocompatibility for NGC. Further, applying electrical stimulation to the MWCNT-PEGDA 3D scaffolds (500 μA and 1 mA) had improved cell viability and neurite extension with an increased neuronal marker (Nestin, MAP2, TUJ1, GFAP) expressions than low-stimulated MWCNT-PEGDA scaffolds (100 μA) [
]. Similarly, rGO-incorporated grid-like PCL scaffolds were fabricated using EHD-jet 3D printing technique with an average fiber diameter of 46±5 μm and pore size of 125±15 μm. The addition of nanostructured rGO (2.5 mg) provides a large surface area-to-volume ratio to enhance cell infiltration and proliferation (cultured with PC12 cells). However, rGO-incorporated PCL-based EHD-jetted scaffolds reduced the mechanical strength (Young's modulus – 16.76±3 MPa; yield strength – 1.51±0.3 MPa) significantly than PCL scaffolds (Young's modulus – 91.25±4 MPa; yield strength – 4.45±0.7 MPa). In contrast, the conductivity of the rGO/PCL scaffold increased up to 1.35±0.3 mS/cm compared to the non-conductive PCL scaffold (0.09±0.005 μS/cm) [
]. These results suggest that while utilizing conductive components in bioinks or biomaterial inks, external electrical stimulation of printed constructs directly influences the neural cell regulatory activities with oriented growth and provide a potential treatment for neural disorders.
3.5 Mechanical property
The peripheral nerves exhibit viscoelastic behavior with unique mechanical characteristics depending on their location and vary between species (e.g., Young's modulus of the sciatic nerve: Rat – 576 ± 160 kPa [
]). Thus, the 3D-printed NGCs should have mechanical properties (such as tensile strength, compressive strength, stiffness, elasticity, kink stability and suture ability) similar to native nerves, in order to prevent damage or compression to the injured nerves and surrounding tissues of the implantation site. Several reports have emphasized the importance and relationship of mechanical properties with other parameters, such as micro-structural design, material composition, and conductivity, as they aid in developing 3D conduits that are mechanically-equivalent to the native nerves. Further, several reports have also emphasized optimizing these properties (e.g., stiffness, tensile strength, bending modulus) as they influence neural cell behavior and orientation. For example, the rat Schwann precursor cell line (SpL201) exhibited better cell attachment and proliferation in stiffer substrates (1–100 MPa) compared to PC12 cells, possibly due to higher expression of integrins, focal adherents and signaling pathways involved in actin−myosin interactions [
]. Dilla et al., synthesized different block length ratios of di-block and tri-block copolymers using monomethyl ether polyethylene glycol (PEG), polypropylene maleate (PPM), and polypropylene fumarate (PPF), represented as PPFmPEGnPPFm polymer. To study the printing feasibility of the synthesized copolymer, a gyroid-patterned 3D scaffold with a strut diameter of about 100 µm was fabricated via continuous DLP technique using the optimized concentration of triblock PPF3PEG23PPF3 copolymer (50%) mixed with 3% bis-acyl phosphine oxide (photoinitiator), 0.4% Irgacure 784, (light scattering agent) and 0.7% oxybenzone (radical scavenger). The PPFmPEGnPPFm copolymers were prepared with different block ratios (PPF3PEG23PPF3/water and PPF3PEG45PPF3/diethyl fumarate) to evaluate the tensile properties after the photo-crosslinking process. Results demonstrated a 10-fold increase in extension at break values for PPFmPEGnPPFm/water hydrogels compared to PPFmPEGnPPFm/diethyl fumarate-based hydrogels, which may attribute to an increase in the crosslinking density due to incorporation of diethyl fumarate (DEF) making it more brittle. In contrast, the water-based PPFmPEGnPPFm resin exhibited lower crosslinking ability, resulting in the conformational relaxation of the PEG chain until failure. Hence, by altering the material composition, the authors have achieved various tensile extension values to match desired tissue of interest. In addition, these hydrogels (PPF3PEG45PPF3, PPF7PEG45PPF7, PPF12PEG45PPF12, PPF3PEG91PPF3, PPF7PEG91PPF7, PPF12PEG91PPF12) when cultured with Schwann cells displayed good cell viability (>95%), denoting the non-toxicity of the synthesized PPFmPEGnPPFm hydrogels and suitability for peripheral nerve regeneration. [
Singh et al., fabricated a photocurable hollow nerve guidance conduit using methacrylated polyglycerol sebacate (PGSm) using stereolithography at dimensions (5 mm length, 700 mm internal diameter, and 350 mm wall thickness) equivalent to rat sciatic nerve. Initially, the mechanical properties were studied by varying the degree of methacrylation (DM), which showed that increased DM value causes increases in the PGSm hydrogels' stiffness values. The compressive strength and suture retention strength analysis of the SLA-printed PGSm conduits with 0.75 DM value (chosen for in vivo implantation) showed an average compressive Young's modulus of 3.2 MPa and average suture retention strength of 12.3 MPa, highlighting the flexibility, also supporting multiple 9–0 polyamide sutures of PGSm NGCs and may be suitable for in vivo implantation purposes [
Yet another important criterion to consider is the degradation ability of the developed PNC when implanted in vivo. Ideally, the degradation rate of a PNC should match the rate of axon growth and ECM formation, strong enough to maintain its shape without any kinking/collapse throughout the regeneration stages, with minimal swelling ability, and cause no inflammatory reactions[
]. When the degradation rate of the conduit is faster, it may lead to swelling with poor mechanical strength due to an increase in porosity, which will eventually form scar tissues surrounding the implant. On the other hand, if the conduit degradation is slow, it may develop compression in the newly formed axons and cause chronic inflammation. In addition, the degradation rate of PNC should be faster in the proximal ends and degrade slower in the distal nerve ends, particularly for long-distance nerve gaps[
Hu et al., have successfully developed hollow and bifurcated nerve guidance conduits (NGCs) using cryo-polymerized gelatin methacryloyl (cryoGelMA) hydrogels integrated with supportive adipose-derived stem cells (ASCs) via indirect 3D printing technology. The pre-hydrogel solution (containing GelMA, ammonium persulfate, and tetramethylethylenediamine) was poured into the printed 3D "lock and key" molds, cryo-polymerized for 24 h at –20°C, followed by removal of the molds to obtain NGCs with an inner diameter of 1.5 mm, an outer diameter of 4.0 mm and length of 15 mm. In vitro degradation of the developed CryoGelMA NGCs in collagenase solution showed complete degradation within 20 h. Later, the acellular hollow conduits were implanted subcutaneously on the dorsal side of the SD rats for 8 weeks, and evaluated the in vivo degradation potential of the fabricated conduits. The conduits remained structurally stable for up to 2 weeks, followed by the collapse of the conduit in the fourth week, but it did not degrade completely until 2 months. The NGCs showed mild foreign body reactions with fibrous tissue and neo-capillary formation in the fourth and eighth weeks. Since the conduits degrade at a proper rate (2-4 months), similar to commercially available NGCs (e.g., Neurotube®, AxoGuardTM Nerve Connector), these conduits can be suitable nerve grafts for peripheral nerve regeneration[
]. In another study, Chen et al., had prepared a photo-crosslinked peripheral nerve conduit using digital light processing (DLP) 3D printer by adding decellularized rat sciatic nerve ECM solution to the mixture of water-based polyurethane (PU) resin and polydopamine (PDA) solution (PU/PDA). The conduits were fabricated at a length of 1.4 mm with inner and outer diameters of 2 mm and 2.5 mm, respectively. The hollow conduits were immersed in simulated body fluid at physiological temperature to determine the in vitro degradation behavior. Results revealed that the residual weight of the conduits had 93%-96% and maintained their structure for up to 12 weeks, which may be sufficient for the native nerves to regenerate completely[
]. The optimized degradation time of the nerve conduits depends on the defect length, as longer nerve gaps require conduits with longer degradation time. For example, a critical-sized nerve gap of 10 mm in rats facilitates axonal regrowth approximately after the third week of conduit implantation. Hence, the developed conduit should initiate the degradation process only after the regeneration of axons in the proximal end of the defect site, which may prevent nerve entrapment symptoms at the implanted site and secondary surgeries to remove the conduit. Clinicians widely use non-degradable conduits such as silicone rubber for treating peripheral nerve injuries due to their bio-inertness and high mechanical strength. Yet, these conduits are poorly degradable and develop tension in the neighboring tissues, requiring secondary surgery for conduit removal. [
Clinical biomaterials or devices should be sterile as they come into contact directly with the tissues. They must follow specific sterilization standards such as BS EN 556-1, ISO 11137, ISO 20857, and ISO 11135 and confirm the sterilization safety along with assurance of sterility for the specified expiry period [
]. There are several methods of sterilization, such as radiation-based (X-ray, electron beam, and gamma), chemical-based (ethylene oxide, hydrogen peroxide or ozone oxidation and gas plasma), and thermal-based (dry heat) methods. Both clinicians and scientists should consider appropriate sterilization techniques critically to ensure that the functionality of the materials or the device is not affected. For instance, some terminal medical devices, which are made up of natural/synthetic materials (e.g., collagen, dECM, PVA), may not be sterilized using certain sterilization methods such as UV light, dry heat, and moist heat sterilization as the radiation or the applied heat may alter the material structure, physicochemical properties, or functionality [
]. For instance, the gamma radiation sterilization method reduced the molecular weight of electrospun PCL fibers (Mn - 17,600 g/mol) compared to other sterilization methods (EtOH – 21,700 g/mol, PAA – 19, 900 g/mol, EtOx – 18,900 g/mol and UV – 19,900 g/mol), which were determined using gel permeation chromatography. Moreover, the morphology of the fibroblasts cultured over the gamma-irradiated PCL fibers for 7 days showed good cell spreading with increased cell density compared to PCL fibers sterilized in ethylene oxide (EtOx), peracetic acid, ethanol, and UV irradiation [
]. In addition, the polyurethane-based biomaterial scaffolds were sterilized efficiently using plasma surface modification using argon gas compared to other methods such as autoclave, gamma irradiation, ethanol, ethylene oxide, UV, chemical bleach by sodium dichloroisocyanurate dihydrate (SDIC) and antibiotic-antimycotic treatment. The scaffolds sterilized using plasma surface modification using argon gas also showed good cytocompatibility compared to other sterilization methods [
]. Hence, the right choice of sterilization technique should be chosen which is ideal for the selected biomaterials. Most FDA-approved peripheral nerve conduits, prepared via conventional scaffold fabrication methods, are sterilized using gamma irradiation (e.g. NeuroflexTM, NeurolacTM), electron beam irradiation (e.g., Versawrap nerve protector, Salumedica nerve cuff), and ethylene oxide sterilization (e.g., Avance nerve graft, NeuroMatrixTM, NeuroFlexTM) [
]. In situ 3D printing/bioprinting is the ultimate achievement of an AM technique, where 3D printers/bioprinters are implemented in the operating rooms and enable the printing of patient-specific tissues/organs instantaneously during emergencies [
]. Unlike 3D bioprinters, most 3D printers are placed in non-sterile rooms, which necessitates complete sterilization of 3D printed parts and examination of alteration in their properties after sterilization [
]. In addition, the bioprinted scaffolds, which are fabricated using biopolymers, cells, small biomolecules, growth factors or enzymes, may not be suitable for common sterilization protocols as temperature or radiation may damage the scaffold characteristics, thus requiring sterile filtration of the scaffold material components using suitable-sized sterile filters or the addition of antibiotics before the printing process to prevent microbial contamination. More specifically, the physicochemical properties, including the rheological properties and printability of the biomaterial inks and cell-laden bioinks, may get affected depending upon the common sterilization process and may eventually affect the mechanical strength and stability of the 3D-printed constructs under physiological conditions [
]. Hodder et al., had prepared a homogeneous mixture of 3% alginate/9% methylcellulose ink and investigated the rheological properties and printability of the ink after subjecting it to various sterilization methods, such as autoclave, supercritical carbon dioxide (scCO2), UV-irradiation and gamma-irradiation. Rheological results displayed shear thinning behavior for both sterilized and non-sterilized Alg/MC samples. However, the viscosity greatly reduced for gamma irradiation-sterilized group (11±1.5 Pa.s) and slightly reduced for the UV-sterilized group (224±12.4 Pa.s) compared to the autoclave-sterilized (291±13.3 Pa.s), scCO2-sterilized (273±27.3 Pa.s) and non-sterilized groups (308±7.5 Pa.s). In addition, the gamma irradiation-sterilized group also showed poor stability and stacking ability during the printing process, requiring an immediate crosslinking process to achieve structural integrity [
Naghieh et al., fabricated 3D-printed scaffolds by impregnating low alginate concentration on the gelatin sacrificial framework. Initially, gelatin was printed to form a 20-layered grid-shaped scaffold with a dimension of 25 × 25 × 2.5 mm and a pore size of 2.5 mm. These 3D-printed gelatin scaffolds were impregnated with alginate solutions (0.5%, 1.5%, and 3.0%), followed by cross-linking with 50 mM CaCl2. Later, the alginate-impregnated scaffolds were incubated at 37°C to completely remove the printed gelatin framework. The fabricated scaffolds were sterilized for 20 minutes by two different sterilization methods, such as ethanol and UV, to understand the effect of sterilization on the mechanical stability of scaffolds and compared with the non-sterile scaffolds. The elastic modulus values obtained from compressive stress-strain analysis of the prepared scaffolds decreased in the order of high-concentrated (3%) ethanol-disinfected alginate scaffolds < low-concentrated (0.5% and 1.5%) ethanol-disinfected alginate scaffolds < UV sterilized scaffolds < non-sterilized scaffolds, which represents ethanol-disinfected scaffolds did not drastically change the mechanical properties of indirectly printed alginate scaffolds. UV-sterilized and non-sterilized scaffolds showed poor mechanical strength due to faster degradation and alteration in the alginate backbone on exposure to UV light and bacterial contamination. Further, high-concentrated (3%) ethanol-disinfected scaffolds also maintained good structural integrity for almost one week and a slow degradation rate compared to 0.5% and 1.5% alginate scaffolds, which may attribute to the presence of high crosslinking density [
Compared to traditional scaffold fabrication methods, the recent advancements in additive manufacturing have tremendously impacted tissue engineering approaches in recapitulating 3D tissue analogs with intricate structures of native tissues embedded with homogenous/heterogeneous cell types, which are eventually tested by in vivo implantation studies (particularly for regenerative medicine applications), disease modeling and drug screening applications [
]. Various attempts have been made to manufacture nerve-equivalent conduits via AM methods that emulate the structure, microenvironment, and physiology of the native nerve tissues and provide successful preclinical outcomes with better functionality. In addition, the 3D printed/bioprinted nerve conduits possess several advantages, including patient-specificity, design flexibility, low cost, and ease of preparation in a minimum printing duration and post-processing steps with high precision, efficacy, and reproducibility. However, manufacturing effective 3D-printed nerve constructs is a big challenge. Some of the challenges are presented below:
Developing novel biomaterial inks / bioinks and crosslinking strategies: Each AM technique requires unique biomaterial-based parameters such as concentration, viscosity, rheology, and crosslinking parameters. Until now, very limited biopolymers such as alginate, gelatin, PEG, PCL, and PU were used as biomaterial inks (repetitively used in various combinations) for developing 3D NGCs and were predominantly crosslinked via ionic (e.g., Ca2+) and photo-crosslinking (e.g., acrylate groups) mechanisms. These crosslinking methods used in AM techniques own remarkable limitations, including poor stability and brittleness when studied for a longer duration, color change in the printed product, highly toxic when used in higher concentrations or molecular mass, etc. In addition, dECM-based hydrogels were used to develop 3D NGCs even in harsh environments (e.g., heat, UV light) as they provide a biomimetic environment with nerve-equivalent mechanical properties. However, these dECM hydrogels restrict their use as a novel biomaterial in developing an ideal NGC, requiring more research on stability, degradation, batch-to-batch variations in the manufacturing process, etc. Hence, a deeper understanding of other natural and synthetic biopolymers, their chemical composition, and crosslinking mechanisms are essential to identify smart bioinks/biomaterial inks suitable for fabricating NGCs with intricate structures, which may synergistically provide appropriate microenvironment and physiological (including physical, chemical, biological and mechanical properties) cues for the regrowth of nerve cells.
Engineering personalized NGCs have not been clinically employed as they require more optimizations related to 3D scanning and defect-free reconstruction of a nerve and printing speed and accuracy of the 3D printer/bioprinter, which involves combined effort from clinicians, biomedical engineers, and computer experts. Though there are several advancements in 3D imaging, reconstruction procedures, and AM techniques, creating an appropriate resolution for personalized nerve constructs is one of the technical hurdles in constructing an ideal NGC. In addition, the implantation of the 3D-printed and bioprinted NGCs is crucial due to surgical challenges involved in properly attaching the printed structures with native nerves, unexpected damage in the printed parts, and sterility issues.
Multi-functionalized NGCs: Treating long-distance nerve gaps (> 3 cm for humans) with commercially available conduits has shown a lesser meaningful recovery rate in both sensory and motor neurons[
] and is thus best suited for short nerve gaps. In addition, these conduits have led to several complications, including infections, pain, stiffness, and amputations, resulting in implant removal. Thus, the uniform-encapsulation of live cells (e.g., stem cells, Schwann cells, patient-derived cells) and gradient-encapsulation of drugs, nerve growth-specific peptides, and growth factors in additively-manufactured NGCs may benefit long-distance nerve injuries, aiding in rapid nerve regrowth and recovery in a limited time without any complications.
Scale-up of 3D nerve constructs: Generally, scaling up the 3D printed / bioprinted tissues involves two approaches - vascularized thicker constructs and in-situ bioprinting. Developing large and functional nerve tissue requires the hierarchical incorporation of heterogeneous cell types, blood vessels, lymphatic vessels, and neural networks. A few attempts have been made to integrate vascular structures within the printed nerve constructs to overcome diffusion limitations and increase the cell survival rate in deeper regions of the constructs. However, differentiation and maturation of printed cells with multi-scale vascularization and functionality may be enhanced by placing the bioprinted or cell-seeded constructs in a suitable bioreactor. Another approach for scaling-up 3D nerve constructs is using in-situ bioprinting using a robotic arm, printing directly (depending on the scanned data of the damaged area) at the defect site to restore the damaged nerves, though successfully employed for calvaria and dermal defects.
Ethical concerns: Though AM techniques reduce the demand for animal testing, applications such as drug screening and testing can be carried out in physiologically equivalent 3D-printed organoids/organ-on-a-chip. However, there are no ethical and regulatory concerns for printable biomaterial inks and their modification, bioinks (e.g., stem cells), 3D printers/bioprinters, storage of materials and printed products, and imaging and reconstructing modalities [
]. Stringent rules and medical practices are required to fabricate human-scale 3D-printed/bioprinted constructs to prevent illegal organ trading.
From the unmet challenges in nerve graft commercialization, the strength, weakness, opportunity, and challenges (SWOC) analysis has been illustrated in Fig. 9. These SWOC must be addressed with long-term efforts, which would pave the way towards creating and commercializing 3D-printed/bioprinted nerve guidance constructs in the near future. In addition, the recent expansion of 3D printing technology, such as four and fifth-dimensional printing, extends 3D printing to another futuristic research approach. However, these strategies were minimally explored in nerve tissue engineering applications, necessitating thorough knowledge generation through extensive research.
Declaration of Competing Interest
The authors declare that there is no conflict of interest.
The authors wish to acknowledge Nano Mission, Department of Science and Technology (DST) (SR/NM/TP-83/2016 (G)), and Prof. T. R. Rajagopalan R & D Cell of SASTRA Deemed University for financial and infrastructural support. We also wish to acknowledge ATGC grant, Department of Biotechnology (DBT) (BT/ATGC/127/SP41147/2021), Adhoc funding, Indian Council of Medical Research (ICMR) (17 × 3/Adhoc/23/2022-ITR) and DST SERB CRG (Exponential Technologies) grant (CRG/2021/007847) for financial support. First author is thankful to Innovation in Science Pursuit for Inspired Research (INSPIRE) scheme of DST, Government of India, for Junior Research Fellowship (IF150843).
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