Pre-bioprinting

Pre-bioprinting involves the development of a printer-useable CAD model and the development of a bioink to be subsequently used for printing. This step conventionally involves technologies like magnetic resonance imaging (MRI) scans and computed tomography (CT) scans which provide images that are then tomographically reconstructed to produce two dimensional (2D) images. Once the design is completed in CAD, it will proceed to the slicing process, which will generate G-code for the transition to the standard tessellation language (STL) file, which can communicate with a printer [57]. This stage also encompasses the selection of biomaterial(s), and the obtaining of cells at desired densities through direct isolation from tissue biopsy, or conventional sub-culturing methods, for bioink development. To keep the cells viable, oxygen and other nutrients are combined with it. The pre-bioprinting phase is crucial in influencing the characteristics of the constructs produced during bioprinting. It is essential to follow the proper methods during the pre-bioprinting stage to guarantee that the required cell quality is obtained which can then promote tissue formation after bioprinting [58].

Bioprinting

This is the stage where the actual printing process occurs. The bioink is then loaded into the printer cartridge, which deposits the substance in accordance with the predetermined digital model developed during the pre-bioprinting stage. To create 3D tissue structures that resemble biological tissues, cell-laden bioinks are deposited in a layer-by-layer manner onto the substrate. Crosslinking is an essential event following successful bioprinting that has a substantial impact on the biomechanical stability and physicochemical properties of bioprinted constructs as well as the cellular behavior of loaded living cells [59]. This part of the bioprinting procedure is challenging since it sometimes necessitates printing using several cell types depending on the tissues and organs being created [58].

Post-bioprinting

Post-bioprinting is the final step in the bioprinting process that is necessary to give the printed structure functionality. In this phase, the tissue structure is developed in an incubator or more preferably in a bioreactor. Physical and chemical stimulations are required to maintain the structure and function. These simulations send signals to cells, triggering them to reconfigure themselves while sustaining tissue growth. Because post-processing conditions have a strong influence on cell-cell interactions, a wide range of stimuli can be used to modulate the tissue maturation process during post-bioprinting processes. The printed construct is subsequently evaluated and biologically compared to native tissue before being used for its intended purpose [58].

Extrusion-based bioprinting

The most popular technique for printing 3D biological structures is extrusion-based bioprinting (EBB), also known as micro-extrusion. Several research groups have employed this bioprinting technology for research involving tissue engineering. EBB is considered the most commonly used bioprinting modality because of its versatility and the wide availability of biopolymers that can be extruded. In essence, EBB is a nozzle-based printing technique in which the bioinks are continuously and precisely extruded through a syringe nozzle during the printing process. The variety of material selection across a wide range of bioink and the high accuracy of production of chemically appropriate tissues or organs are the most promising aspects of EBB. The shear-thinning biopolymers, which aid in extruding the bioink, are most frequently used to create extrudable bioinks. The hydrogels, also known as printable biopolymers, are non-Newtonian fluids in which viscosity decreases under shear stress. The mechanical pressure used in extrusion straightens the entangled polymer chains, lowers the viscosity of the bioink inside the nozzle, and promotes the survival of living cells. Due to negligible shear during printing, the hydrogel regains its viscosity, providing printing fidelity to the finished construct. An extrusion bioprinter may use pneumatic, piston, or screw-based mechanical pressure, depending on the parameters of the hydrogel and the cells. The resolution, precision, and overall structural fidelity of the printing construct are significantly impacted by the extrusion pressure, nozzle diameter, and printing speed of the extrusion bioprinter. The primary process parameter of the printer for a particular hydrogel is nozzle diameter optimization. High extrusion pressure is necessary for a nozzle with a small diameter for a viscous hydrogel, and this pressure also increases the shear stress on the cells. Higher shear stresses can cause damage to the cells, which lowers the vitality of the cells after printing [60, 61]. The ability to print very viscous hydrogels with high cell concentration at a moderate speed is one of the main advantages of this technique. Additionally, a complex functioning organ can be printed utilizing a multi-head or co-axial head extrusion bioprinter, which allows the printing of numerous cell types using the same or distinct hydrogels within the same construct [62]. Without doubt, EBB is the most preferred technique when bioprinting bioartificial pancreas. Likely so majority of the studies reviewed in this article have been carried out using EBB and its variations to achieve precise bioprinted structures with high shape fidelity and complexity quite close to their native counterparts.

Inkjet-based bioprinting

Another popular technology is inkjet bioprinting, also known as drop-on-demand bioprinting, and it can be used for both biological and non-biological purposes. The ink in the cartridge was eventually changed to biological material, and the paper was replaced with an electronically controlled elevator stage to give control. Initially, this technology was exclusively utilised for 2D ink-based printing. At the moment, inkjet bioprinting may be carried out on bioprinters specially created to handle and print biological materials with great precision, speed, and resolution. This technique is based on drop-on-demand techniques. Different biosimilars are encased within the hydrogel matrix, and premixed bioink solutions are placed inside the inkjet cartridge to be depleted dropwise through the inkjet printhead in a regulated manner. The bioink chamber can produce the bioink droplets by applying either thermal energy [63, 64] or piezoelectric impulses [65, 66]. Tiny air bubbles created by thermal energy during thermal inject bioprinting create pressure pulses within the bioink solution that force out bioink droplets of various diameters. The temperature gradient, the viscosity of the bioink, and cell concentration all affect the size of the droplets. For piezoelectric inkjet bioprinting, a polycrystalline piezoelectric ceramic material transforms the applied electric current to transient mechanical pressure, causing a droplet of bioink to expel onto a construction platform. Both single and multiple piezoelectric printheads are possible. Different cell kinds can be deposited in the same printed construct at the same time using other printheads [67]. Piezoelectric-based printheads can control droplet volume more precisely than thermal printheads, making them more practical sometimes. However, because the typical operating frequency of a piezoelectric printhead is between 15 kHz and 25 kHz, which might harm the cell membrane, many researchers favour thermal inkjet bioprinting over piezoelectric technology [63]. Inkjet bioprinting has a high printing resolution of approximately 50–300 μm [68]. Nevertheless, this method takes a long time since submicrometer-sized droplet volumes are tiny. The main disadvantage of this method is that bioprinting is challenging when the viscosity of hydrogel is high (> 10 cP) and the cell concentration is greater than 5 × 10⁶ cells/mL [69]. When the cell concentration is high, cell aggregation or sedimentation might happen in the nozzle and cause cell clogging. The hydrogels must also be sufficiently wettable and have the right amount of surface tension to pass through the cartridge and nozzle. Additionally, the printed structures obtained by this method exhibit weak mechanical integrity due to low viscosity. These factors have limited the widespread usage of this technique for the development of bioartificial pancreas.

Laser-assisted bioprinting

The process of depositing biomaterials onto a surface while employing a laser as an energy source is known as laser-assisted bioprinting (LAB). This method had previously only been used to transfer metals, but it has recently been altered to work with biological components like cells, DNA, and peptides. A layer of biological material generated in a liquid solution with a receiving substrate facing the projector, along with a ribbon with donor transport support, makes up a laser-assisted bioprinter. In LAB, biomaterials such as hydrogel, culture media, cells, proteins, and ceramic materials will be employed. LAB is a nozzle-free dispensing technique that allows for a wide range of bioink viscosity and cell concentration changes without clogging the nozzle. While the cell concentration can be increased up to 1 × 10⁸ cells/mL without causing nozzle blockage, the bioink viscosity can be adjusted between 1 mPa·s and 300 mPa·s without impacting the printing resolution [69]. A ribbon, commonly a laser-transparent material made of either glass slide or quartz, is part of the printer arrangement. The donor side of the ribbon is coated with a laser-absorbing medium (such as Ag, Au, Ti, or TiO₂) before the cell-encapsulated hydrogels are sprayed on top of the laser-absorbing coating. The absorbing media are focused by the laser impulse through the ribbon, and as they evaporate, they raise the local pressure on the bioink film. In the direction of the bioink film, cavitation-like bubbles are produced by the absorbing media’s vapour pressure. The bioink layer’s jet is formed by the bubble’s expansion and contraction, which causes the bioink droplets to be created and transmitted to the printing substrate [70, 71]. The LAB has numerous benefits, but there are also some downsides, such as the constant contamination of the printed substrate by the absorbing media. The fragile spray coating used in LAB can quickly dry on the ribbon surface before printing [72]. These bioprinters move at a medium speed, and roughly 95% of the cells are still viable after the process. However, random cell distribution results in cells spreading onto the ribbon, which results in uneven cell printing. The major limiting factors towards the widespread implementation of LAB for bioprinting include involvement of high costs and limited materials being reported which can be printed using this technique.

Main cells

One of the most challenging situations for researchers is the selection of source and type of cells while bioprinting them to replicate islet function and expedite neotissue formation. While there is an apprehension towards using allogenic and xenogenic sources due to their inherent immunogenic response, obtaining autologous cells may often be difficult and associated with donor-site morbidity. From a pancreas bioprinting perspective primary islets derived from allogenic or xenogenic sources have been encapsulated during bioprinting to develop pancreatic tissue. Primary islets are often recognized as the preferred cells since they are the native cells forming the pancreas, and can be obtained through a minor biopsy of the pancreas to subsequently extract islet cells. Although isolated islets are a common source for pancreas bioprinting, they have significant limitations including an additional surgical procedure to harvest them causing donor site morbidity, limited growth, and loss of insulin-producing capability during in vitro culture, are difficult to expand during culturing, and thus have low intrinsic healing capacity [86]. Basically, when islets are isolated, the ECM and islet vasculature are destroyed, which may have a deleterious impact on islet function after transplantation.

An alternative to the issues associated with primary cells is the use of immortalized cell lines closely resembling native islet function. They offer several advantages, such as they are cost-effective, robust, easy to use, providing an unlimited supply of cell sources, and bypassing ethical concerns associated with the use of animal and human primary cells. Insulinoma cell lines such as MIN6 [87], and INSE-1 [88], BRIN-BD11 [89] have been successfully used in bioprinting applications for replicating native islet function. Unlike primary cells, immortalized cell lines can replicate forever and are typically more robust and user-friendly. However, working with cell lines has a number of drawbacks, including the fact that they are genetically engineered. Moreover, variability in cultures can be brought about by genetic drift or extensive passaging of cell lines, which can lead to genotypic and phenotypic heterogeneity over time [90]. Cell lines may therefore not accurately represent primary cells and may produce inconsistent findings.

Another alternative to primary cells, induced pluripotent stem cells (iPSCs), which are obtained through cellular reprogramming of somatic cells have been investigated for their ability to form islet cells [91]. Pluripotent stem cells are the progenitors of all adult tissues throughout development. Technological improvements in differentiation protocols have facilitated the robust development of pancreatic β-cells from patient-derived iPSCs. Differentiation protocols for iPSC-derived islets primarily target the same developmental stages, beginning with the definitive endoderm and primitive gut tube, then narrowing cell fate to pancreatic and endocrine progenitors, before ultimately targeting the final differentiated β-cell. Growth factors are commonly used to modulate the pathways required for differentiation stages, with the goal of mimicking embryonic development [91]. Over the last decade, β-cell and pancreatic progenitor stem cell differentiation protocols have been meticulously developed, leading to functional iPSC-derived islets. Due to their non-invasive collection from the patient and autograft without immunomodulation, human iPSCs (hiPSCs) can be thought of as an advantageous substitute for stem cells used in regenerative medicine, without ethical considerations as in the case of applications involving embryonic stem cells (ESCs). Regardless of the fact that iPSCs have very low yield and are extremely immature, their in vitro differentiation into islets is quite challenging, in broad sense, there are technical challenges and financial concerns associated with iPSCs. Chromosome aberrations or mutations are common when reprogramming somatic cells to iPSCs. Furthermore, the formation of teratomas remains a concern when using iPSCs. Even though advances in iPSC-derived islet differentiation protocols are promising, a universal differentiation protocol that could be applied to all iPSC lines would make autologous cell therapy more feasible [91].

Adult stem cells (ASCs), on the other hand, represent a dependable alternative cell source. Compared to stem cells belonging to embryonic origin that are able to exhibit pluripotency, ASCs are multipotent and lineage-committed cells having a more limited capacity to differentiate into the various cell lineages of our body. However, there are ethical considerations regarding the use of ESCs in research due to the inevitable destruction of human embryos, thereby shifting the focus on ASCs [92]. ASCs reside in distinct niches with microenvironments that enable them to maintain an undifferentiated state and to replace specialized cells in the affected tissues that have been damaged [92]. Mesenchymal stem cells (MSCs), which are spindle-shaped cells initially misidentified as fibroblasts, are among the most widely used ASCs. The use of MSCs in tissue engineering can be supported by their active role in tissue regeneration and production of cells as a replacement for damaged and dead cells [93]. Recent studies have shown that MSCs are able to migrate to distant areas where damage has occurred and potentially offer reparative cells or produce soluble trophic factors through paracrine signalling which aids in cell survival, cell proliferation, and cell migration to augment tissue regrowth [94]. Furthermore, MSCs have anti-inflammatory and immunomodulatory properties, which enable them to reduce inflammation and restore or suppress immune cell functioning [95]. Most importantly, prior to usage in tissue engineering applications, MSCs are able to be readily cultured in vitro to optimal cell numbers over multiple passages without losing their self-renewal ability [96]. Given the functionalities listed above, the use of MSCs in pancreatic tissue engineering approaches has the potential to overcome the long-standing problem of finding a suitable cell source. Overall, MSCs are superior to other forms of stem cells in terms of ease of isolation, plasticity, and clinical translation to generate autologous cells. Although there exists no standard protocol for obtaining pancreatic β-cells from MSCs, multiple differentiation protocols reported by various research groups mention the use of a cocktail of small molecules and growth factors to guide MSC-differentiation towards pancreatic β-cells [97]. Although providing insight into these protocols is not a part of this review, the authors recommend recent reviews by Pavathuparambil Abdul Manaph et al. [97] and Wszoła et al. [98] for a comprehensive understanding of the topic.

Supporting cells

Technological advances in 3D bioprinting offer the use of multiple nozzles which enables the incorporation of different cell types at designated niches within a bioprinted scaffold to replicate the high complexity of tissues. Most often, the use of supporting cells alongside main tissue forming cells assists better neotissue development and, hence function. The most important factor influencing the effectiveness of islet transplantation is hypoxia [99]. This presents a major issue in the case of immunoprotective layer formation techniques such as encapsulation which is employed in bioprinting. Since 3D bioprinting primarily produces macroscale capsules, the issue of extravascular macroencapsulation, such as core hypoxia, happens since the cells are isolated from their surrounding blood vessels. As a result, the viability of the transplanted islets in the subcutaneous site is reduced significantly. Within the bioprinted constructs, the formation of a vascular system can be induced through the addition of supporting cells to create a relatively homogeneous blood flow. The islet of Langerhans is densely vascularized with fenestrated endothelium, which allows cells to detect blood glucose and secrete insulin into the systemic circulation [100]. Endothelial cells are also important in promoting islet function by upregulation of insulin transcription via cell-to-cell contact or humoral factor secretion [101]. Co-culture with endothelial progenitor cells, or human umbilical vein-derived endothelial cells (HUVECs), represents a promising strategy to promote vascularization within bioprinted constructs. Differentiating into endothelial cells, the building blocks of blood vessel endothelial lining these cells can undergo crosstalk with islet cells and promote insulin expression and secretion, as well as islet survival within bioprinted constructs to form fully functional biomimetic pancreatic tissue [102].

As part of the alloresponse following implantation of bioprinted tissue, recipient T-cells recognize alloantigens of the allograft, which activates recipient T-cells to evoke an inflammatory response at the implant site [103]. Such a response causes acute graft rejection in addition to chronic graft dysfunction, necessitating the administration of immunosuppressive drugs to prevent the onset of alloresponse. However, immunosuppressants are associated with negative side effects, therefore mandating the need to develop novel immunotherapies to achieve immunosuppression-free transplantation. T-cell sub-populations known as “regulatory T-cells (T_regs)” are experts in controlling and suppressing the immune system. Previously, in mouse hepatic allograft models, it has been demonstrated that T_regs play a significant role in spontaneous graft tolerance [104, 105], and there is an increase in T_regs proportion in patients who are spontaneously tolerant [106]. As a result, numerous strategies for making use of their innate abilities have been thoroughly researched. Particular benefits of adoptive transfer of ex vivo expanded T_regs include improved control over the creation and expansion of T_regs, the ability to conduct functional and phenotypic analyses prior to delivery, and more precise control over dosage and delivery timing [107]. This has shown promise in murine islet transplantation models [108, 109].

Alginate

Alginate is a naturally occurring polymeric material derived from the cell wall of brown algae. It is composed of D-mannuronic acid (M-block) and L-guluronic acid (G-block) linked through a 1,4-glycosidic bond [121]. Being an anionic polysaccharide-based biopolymer, it can be easily crosslinked using a divalent cation Ca²⁺, Sr²⁺, or Ba²⁺ enhancing viscoelastic and water-retention properties along with its increased cell binding affinity [122]. This makes alginate an ideal polymer for encapsulating various cell types for bioprinting applications [121, 123]. Alginate has been frequently employed for pancreatic islet encapsulation due to its ability to achieve immunological isolation. Alginate-encapsulated pancreatic islets have been successfully transplanted into people [124]. Many researchers concentrated on alginate and alginate-based hydrogel blends as coatings to produce immune-isolating encapsulation since these polymers have been proven to be non-immunogenic if suitably purified [123, 125]. Pancreatic islets have been observed to sustain their viability and functionality for a while after either macroencapsulation or microencapsulation in alginate was carried out [126]. However, capsule size and composition unquestionably affect islet function, primarily as a function of diffusion distances [127, 128]. The single islet that makes up each microcapsule is enclosed in a thin layer of polymer, reducing the distance between the islet and its surroundings while maintaining the barrier’s immunoprotective properties [129]. However, while macrocapsules may typically be recovered with ease in cases of graft failure, it is virtually difficult to recover microcapsules. Alginate has been frequently employed for pancreatic islet encapsulation due to its ability to achieve immunological isolation. However, as demonstrated in vivo, alginate is vulnerable to losing structural stability and integrity as a result of ion exchange with the environment [130]. Additionally, alginate lacks cell adhesion RGD motifs as well, which reduces cell survival due to anoikis [131].

The earliest attempt at bioprinting the complex micro-architecture of the pancreas was reported in 2010 by Xu et al. [132] using alginate/gelatin/fibrin hydrogels to encapsulate adipose-derived stem cells (ADSCs). The chemical similarity of gelatin and alginate to the ECM, while the property of fibrin regulates cell differentiation and self-organization into a functional tissue structure [133, 134]. Subsequently, when pancreatic islets were introduced at designated micro-holes within these ADSC-laden alginate/gelatin/fibrin hydrogels, the ADSCs were induced to differentiate into vascular endothelial cells and adipocytes to mimic the complex tissue structure of native pancreas. The resulting bioprinted constructs were stable for up to 8 weeks, and facilitated cell growth, organization, and differentiation. Following incubation of constructs in epidermal growth factor (EGF) induction of vasculature was confirmed by positive immunostaining for CD31 and CD34. Alongside, Oil Red O staining carried out confirmed the differentiation of ADSCs into spherical adipocytes. Assessment of whether the bioprinted constructs were capable of mimicking in vivo function was carried out by measuring dynamic insulin release patterns in response to 15 mM glucose stimulation, showing significantly higher insulin secretion in assembled bioprinted constructs compared to the control group comprising only free β-cells. These results were also supported by glucose consumption analysis, where seeding β-cells in the 3D structure significantly increased glucose consumption. This study is a significant achievement in the field of complex organ engineering, which has shown great promise in the establishment of physiological pancreatic tissue.

Alginate because of its excellent biocompatibility and lower toxicity has led to testing in a variety of 3D bioprinting applications [135]. Alginate has a favourable effect on cells, however because of its low viscosity, printing fidelity is compromised, making it difficult for bioprinted scaffolds to retain their shape after being deposited from the printer nozzle [136]. To address this issue, Duin et al. [137] encapsulated islets isolated from rat within an alginate/methylcellulose hydrogel blend to fabricate macroporous hydrogel structures via 3D bioprinting (Figure 4). The resultant bioink was printable and had a higher shape fidelity in comparison to plain alginate. Following the printing process, it was demonstrated via Live/Dead staining and 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) assay that the islets encapsulated within the hydrogel had good viability and metabolic activity. Additionally, the ability of the islets to be able to continuously produce insulin and glucagon in response to glucose stimulation throughout the post-printing stage was reported through dithizone (DTZ) staining and immunofluorescence staining. Through this study it was demonstrated that the alginate/methylcellulose hydrogel does not interfere with the diffusion of relevant macromolecules. Additionally, neither the inclusion of methylcellulose nor the bioprinting process harms the morphology or survival of the encapsulated islets, thereby serving as proof-of-concept that it is possible to 3D plot pancreatic islets that remain functional and respond to glucose stimulation.

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Figure 4. 

3D bioplotting islets in alginate/methylcellulose hydrogel blends. (A) Images of plotted and Sr²⁺ crosslinked hydrogel scaffolds demonstrating high shape fidelity and stability over 21 days in culture. Scale bars depict 5 mm; (B) live/dead staining of free islets representing the control group (left) and encapsulated islets within hydrogel scaffolds (right). Live and dead cells are shown in green and red, respectively. Scale bars depict 200 μm; (C) semi-quantitative assessment of islet viability on the basis of live/dead staining as shown in (B), n > 60 islets; (D) islets stained with MTT on day 1 after bioprinting to check for their metabolic activity (top) and islets stained for presence of insulin with DTZ on day 7 after bioprinting (bottom). Scale bars on the left depict 5 mm; scale bars on the right 2 mm; (E) immunofluorescence micrographs of encapsulated islets stained for nuclei using 4’,6-diamidino-2-phenylindole (DAPI), insulin, and glucagon incubated for 1, 4, or 7 days under cell culture conditions. Scale bars depict 50 μm

Note. Adapted with permission from “3D bioprinting of functional islets of Langerhans in an alginate/methylcellulose hydrogel blend,” by Duin S, Schütz K, Ahlfeld T, Lehmann S, Lode A, Ludwig B, et al. Adv Healthc Mater. 2019;8:1801631 (https://onlinelibrary.wiley.com/doi/10.1002/adhm.201801631). © 2019 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim.

Recently Hu et al. [87] developed a novel bioink by adding the polymer Pluronic F127 to the alginate solution, resulting in improved the printability and flexibility of the crosslinked structure. Meanwhile, to enhance resistance against inflammatory stresses, hypomethylated pectin was added to the bioink composition without changing the viscoelastic property of the bioink. The results of the experiments demonstrated that cellular constructs printed with alginate-pluronic-pectin bioink could allow oxygen diffusion, maintain higher cell viability during exposure to the diabetogenic combination of the cytokines interleukein-1β (IL-1β), interferon-γ (IFN-γ) and tumour necrosis factor-α (TNF-α), and reduce tissue rejections by inhibiting toll-like receptor 2 (TLR2)/1 and ensure the survival of insulin-producing MIN6 cells under inflammatory stress. Moreover, histological examination of the constructs post 4 weeks since their subcutaneous implantation in mice to investigate the tissue response to these constructs in vivo, did not show infiltration of multinucleated giant cells or granulocyte invasion underscoring the biocompatibility of the constructs. This study offers a better engraftment strategy for implanted islets in the treatment of type 1 diabetes.

Gelatin

Another polymer which has extensively been researched as a scaffolding biomaterial for tissue engineering applications is gelatin. It is obtained by the partial hydrolysis of collagen which is abundantly present in the ECM of living tissues [138]. Gelatin is essentially demonstrated thermoresponsive aqueous behaviour at temperature above 30℃ whereas it shows sol-to-gel transition at temperatures below 30℃ making it a suitable scaffold for various bioprinting applications [139]. Additionally, gelatin is biodegradable and biocompatible, with additional properties such as low immunogenicity. Moreover, gelatin is composed of different amino acids, particularly the biologically active RGD tripeptide sequence which promotes cell adhesion by interacting with integrins expressed on cell membranes. Therefore, the addition of gelatin can favor the adhesion, proliferation, and differentiation of beta cells in bioprinted constructs. Several research groups have demonstrated the efficacy of gelatin and its derivatives in encapsulating islets. Considering collagen is an essential component of the basement membrane of ECM in the adult human pancreas, gelatin has an advantage over other polymers in terms of providing the cells with a microenvironment close to their native one [140]. Mechanical strength is one of the most crucial properties demanded by pancreatic tissue engineering. This prerequisite is not met by gelatin alone. As a result, numerous gelatin-polymer mixtures have been investigated.

Another very important consideration while bioprinting structures for islet encapsulation is to maintain optimum porosity within the constructs to enable diffusion of nutrients. It is widely considered that in contrast to conventional bulk hydrogels, islets encapsulated within a porous bioplotted structure could have superior oxygen, glucose, and insulin exchange from surroundings due to a higher surface-to-volume ratio. The proof of concept for bioplotting cells within bioprinted constructs was reported by Marchioli et al. [136] who encapsulated INS1E β-cell line, primary human and mouse islets within an alginate/gelatin bioink. The resulting bioink had printable viscosity and thus could be used for printing macroporous self-standing constructs. Additionally, there were no detrimental effects of the printing process nor cytotoxic effect of material on the cells as validated by their high viability (91%) after 3 weeks. Nonetheless, the metabolic activity of cells encapsulated cells in the hydrogel decreased. The findings of the study revealed that the deterioration of functionality of encapsulated islets is not due to the bioplotting process, but rather results from reduced glucose diffusion through such a viscous hydrogel, such that even a macroporous structure like the one fabricated here cannot restore nutrition transport to reverse this loss of metabolic activity. This study perfectly highlights the inverse relation between the viscosity of material and bioplotting process where a lower material viscosity is beneficial for nutrient diffusion but not printable, while a printable material having higher viscosity creates a barrier for nutrient diffusion and hypoxic conditions affecting cell activity.

Previously in order to fabricate perfusable vascular constructs, gelatin methacryloyl (GelMA), a well-accepted bioink component, has also been used to enhance the printability of alginate/poly (-ethylene glycol)-tetra-acrylate (PEGTA) blended bioink to develop hollow perfusable constructs via coaxial bioprinting [141]. Following this trajectory, in a more recent study Liu et al. [142] co-printed islets, endothelial progenitor cells, and T_regs by adding GelMA to alginate bioink formulations to increase their printability, and stability, and to create an environment that supports transplanted islets in order to maximize islet survival (Figure 5). Scalable core-shell macroporous constructs for islet encapsulation were developed using a coaxial 3D bioprinter. The GelMA/alginate bioink formulations demonstrated shear-thinning behaviour, besides systematically improving to maintain structural stability and cell viability after the bioprinting process. Insulin secretion assessed under low and high glucose settings to assess the islet function following printing, revealed comparable insulin secretion levels in both printed and free islets, suggesting that the method of islet encapsulation and coaxial printing had no negative impact on their survival. Nevertheless, this novel study was the first of its kind to addressing issues related to re-vascularization and immunoisolation of islets via the fabrication of core-shell structures. The coaxial printed construct may shield the encapsulated islets in the core while simultaneously delivering various supportive cells into the shell owing to precise control over the positioning of multiple cells.

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Figure 5. 

Coaxial 3D printed alginate/GelMA scaffolds for islet encapsulation. (A) Schematic representation of the coaxial printing approach carried out to fabricate vascularized bioartificial pancreatic constructs (left), microscope image of coaxial structure of printing nozzle (top right), and bright field image of coaxial strand printed with blue/red dye to visualize core-shell structure (bottom right). Scale bars depict 500 μm (top right), and 2 mm (bottom right); (B) shear thinning behaviour of different formulations of alginate/GelMA bioink. Data are presented in terms of mean ± SD (n = 3); (C) a coaxial printed construct with encapsulated islets and EPCs (left), printed islets with EPC (middle), and islets in core (right). Scale bars represent 200 μm in (B) and (C); (D) mouse pancreatic islets directly after isolation (top left) and immediately after encapsulation in alginate/GelMA (top right); live and dead staining of islets 24 h after encapsulation in alginate/GelMA scaffolds (middle and bottom row). Fluorescein diacetate/propium iodide was used to stain live (green) and dead (red) cells respectively. Scale bars represent 500 μm in (top row), and 200 μm in (middle and bottom row); (E) total secreted insulin after GSIS was measured at day 1 (left) and day 3 (middle) after printing, at low glucose (LG) and high glucose (HG) condition; and secretion ratio of insulin (right) per sample. GSIS: glucose-stimulated insulin secretion; SD: standard deviation; EPCs: endothelial progenitor cells

Note. Adapted with permission from “Development of a coaxial 3D printing platform for biofabrication of implantable islet-containing constructs,” by Liu X, Carter SD, Renes MJ, Kim J, Rojas-Canales DM, Penko D, et al. Adv Healthc Mater. 2019;8:e1801181 (https://onlinelibrary.wiley.com/doi/10.1002/adhm.201801181). © 2019 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim.

pdECM

High post-transplant islet loss appears to be caused by multiple factors, and is a major bottleneck towards their use for therapeutic applications [143]. One direct result of separation from the donor pancreas is the loss of connection between the islets and ECM molecules as well as the disintegration of the extensive vascular network that connects the islets during cell isolation protocols [144]. In addition, the loss of the thick microvasculature is responsible for impairments preventing the pancreatic islets from quickly detecting variations in blood sugar and responding by secreting insulin. These impairments include efficient oxygenation, nutrition delivery, and metabolic waste elimination [145]. Even while transplanted islets can endure for up to several days in an avascular environment, prolonged delays in tissue vascularization cause necrosis and loss of endocrine function [144]. As a result, it is believed that recapitulation of the native pdECM microenvironment, and a thick vascular network are essential for raising the likelihood of successful therapeutic results following islet transplantation procedures. The ECM in native tissue exhibits greater structural, chemical, biological, and mechanical complexity than these materials do [146]. For islets to function better, microenvironmental factors such as interactions between islets and ECM are crucial signals. Collagen, glycosaminoglycans (GAGs), and glycoproteins are some of the several ECM components that can be recreated using decellularized ECM (dECM). For instance, primary islets with peripheral ECMs still present (such as laminin, type I, III, IV, V, and VI collagen, and fibronectin) show lower levels of apoptosis and better insulin sensitivity, suggesting that tissue-specific cell-matrix interactions are crucial for improving the ability of these organoids to function similarly to native tissue [147]. Likely so, organs created through 3D bioprinting should be customized using particular tissue-specific biological inks. This can be done by 3D printing pancreatic islets using a specialized pdECM. Decellularized scaffolds mimic in vivo microenvironment for replanted cells by maintaining the general structure of natural ECM and inherent growth factors [148]. The pdECM hydrogels show improved and better function as derivatives of pancreatic decellularization. According to the concentration of cells and the requirements of the graft, different types of cells can be added to pdECM hydrogels, allowing for modulation of the hydrogel size and pdECM concentration [149]. Therefore, the selection of bioink is crucial in order to accurately recreate the target tissue’s microenvironment in the 3D-engineered tissue construct.

Tissue-specific dECM bioinks are useful for fostering stem cell differentiation and proliferation [150]. Because of its tissue-specific composition and architecture, dECM can act as a cell-favorable habitat in comparison to synthetic polymers [151]. The substantial retention of important components in the dECM should therefore require serious consideration of the decellularization process. In this regard, Kim et al. [152] used pdECM to create a native microenvironment for developing 3D pancreatic tissues that could be transplanted. The primary islets cultured in pdECM bioinks demonstrated excellent viability, enhanced insulin secretion, and functionality. Additionally, co-culture with HUVECs significantly prevented apoptosis of encapsulated islets in the center via the prevention of hypoxia arising from lack of vascularization. The viability of pancreatic islets printed in pdECM hydrogel was comparable to that of the control group comprising of non-printed islets in 3D culture.

Recently, Hwang et al. [153] aimed to create a novel engineering approach for the fabrication of a hybrid encapsulation system that encapsulates aggregates that resemble pancreatic islets using a multi-head bioprinting system and is composed of an assembled macroporous polymer capsule construct and nanoporous pdECM hydrogel. The outer part (macroporous PCL capsule) was fabricated with an interpenetrating architecture, allowing insulin-producing β-cells encapsulated within the hybrid scaffold to maintain their cellular behaviour, such as viability, cell proliferation, and insulin secretion. The inner part (nanoporous dECM hydrogel) composed of 3D biofabricated pancreatic islet-like aggregates was placed into the macroporous polymer capsule. With regard to M1 macrophage polarization which is indicative of inflammation, the developed hybrid encapsulation system demonstrated biocompatibility both in vitro and in vivo. Additionally, each component of the macroporous polymer capsule contained a porous membrane with a sub-micron size enabling the diffusion of nutrients and oxygen to the cells. Additionally, the macroporous polymer capsule offered retrievability and mechanical protection. In conclusion, these findings showed that the 3D bioprinting method makes it easier to create a hybrid islet encapsulation system using materials resembling native microenvironment of tissue may enhance clinical outcomes by promoting the structural maturation and functional advancement of cells.

Another recent study by Idaszek et al. [154] reported development of porous vascularized pancreas grafts via a microfluidic 3D bioprinting platform coupled with a co-axial needle system, using two different tissue-specific bioink compositions: an alginate/pECM bioink encapsulating pancreatic islets, and an alginate/fibrinogen bioink encapsulating HUVECs to develop pancreatic and vascular structures respectively. The resultant bioinks despite having varying rheological properties, were 3D-printed with high shape fidelity and viability thereby promoting endocrine function as reported by insulin secretion in response to glucose stimulation. Additionally, the ability of HUVECs to promote vascularization was confirmed by the formation of intercellular junctions and vessel-like structures by CD31-positive immunostained cells. Overall, the presented findings represented a significant advancement both in the field of biofabrication of vascularized pancreatic tissues and the applicability of multi-material/multi-cellular 3D bioprinting for engineering complex tissues like the pancreas.

Hyaluronic acid

Among the many bioinks currently being studied, hyaluronic acid, and its derivatives stand out because of their physiological relevance, cytocompatibility, shear-thinning capabilities, and ability to be tailored to specific characteristics through chemical modifications. The primary component of ECM and a natural polysaccharide known as hyaluronic acid is employed extensively in regenerative medicine and tissue engineering [155, 156]. It is essentially a non-sulphated disaccharide polymer composed of alternating N-acetylglucosamine and D-glucuronic acid disaccharide units linked by β-1,3 and β-1,4 glycosidic bonds [157]. It is biocompatible and biodegradable, making it completely safe bioprinting applications. According to earlier research, hyaluronic acid added to alginate for islet transplantation improves encapsulated cells’ survival rates [158] while also reducing the body’s immunological inflammatory response [159]. In the realm of tissue engineering, hyaluronic acid has been increasingly utilized as a scaffolding biopolymer. Islet cells that were encapsulated in hyaluronic acid and alginate matrices have reportedly shown high viability and insulin secretion [160]. Additionally, hyaluronic acid-alginate hybrid microcapsules have been demonstrated by Cañibano-Hernández et al. [158] to decrease apoptosis and boost β-cell survival. There is growing evidence available now which suggests that hyaluronic acid maybe a more suitable polymer for encapsulating islets in comparison to alginate which is most often used for cell encapsulation due to lesser immunogenicity.

Hyaluronic acid in its pure form, cannot be utilised as a material in bioink development since it lacks printability and is unstable due to its high-water absorption. Hyaluronic acid methacrylate (HAMA), obtained by the methacrylation process of hyaluronic acid can undergo quick gelation via photo crosslinking with lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP) in the presence of ultraviolet (UV) radiation [161]. HAMA, a novel hydrogel ink for 3D printing, demonstrates quick photosensitivity, rapid gelation, and stable hydrogel performance. HAMA is additionally employed in the creation of biological organoids [162]. As a result, the mixture of HAMA and pdECM could serve as a source of bioink for 3D printing for making islet organoids. Recently, Wang et al. [163] fabricated an islet organoid via 3D printing using a hybrid bioink combining the mechanical properties of HAMA and tissue specificity of pdECM to simulate the in vivo pancreatic microenvironment (Figure 6). The developed HAMA/pdECM hydrogels were reportedly able to regulate islet adhesion, and morphology in vitro through expression of E-cadherin, N-cadherin, and fibronectin, which aids in islet function and activity through the ras-related C3 botulinum toxin substrate 1 (Rac1)/Rho-associated protein kinase (ROCK)/myosin light chain kinase (MLCK) signalling pathway. Consequently, a significant increase in the expression of pancreatic marker genes such as insulin (Ins2) and glucagon (Gcg) were reported via qPCR assay. The HAMA/pdECM hydrogels were consequently used for 3D-printing pancreatic islet organoids via a stereolithographic process, with reportedly high viability of islets observed with HAMA/pdECM thereby underscoring the cytocompatible nature of the material. Similar observations were reported during immunofluorescence analysis which indicated higher insulin and glucagon expressions of islets in the 3D-printed HAMA/pdECM hydrogels. The organoids were moreover able to respond appropriately to glucose stimulation via insulin release and were considered for implantation within diabetic mouse. The 3D printed islet organoids demonstrated excellent robustness in that they were able to retain their functionality even after 12 weeks of implantation, and reportedly restored blood glucose levels within a normal range (6.5 mM/mL) within 60 min of food intake. Additionally, the HAMA/pdECM hydrogels were found to significantly promote re-vascularization in vivo after 90 days of implantation as seen through high density of positive staining for CD31. All these findings demonstrated mimicking the native functionality of islets by the 3D printed HAMA/pdECM islet organoids.

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Figure 6. 

3D-printed HAMA/pdECM pancreatic islet organoids. (A) digital camera micrographs of HAMA/pdECM hydrogels prepared using different concentrations of pdECM; (B) Western blot analysis to detect the protein expressions of E-cadherin, N-cadherin, β-catenin, Rac1, ROCK, and MLCK from HAMA/pdECM hydrogels; (C) quantitative PCR test evaluating the expression of the Ins2, and Gcg in the islets encapsulated within the hydrogels (n = 3); (D) live- and dead-cell staining of pancreatic organoids to detect the effect of 3D printing on the activity of islet organoids (Scale bar = 50 μm); (E) immunostaining micrographs of HAMA/pdECM islet organoids detecting the effect of 3D printing on islet function, including insulin (red) and glucagon (green) (scale bar = 50 μm); (F) glucose-stimulated insulin-release curve from the control, HAMA, and HAMA/pdECM pancreatic islet organoids; (G) blood glucose level in diabetic mice receiving transplantation (n = 6); (H) representative CD31 immunostained images on the surface and cavities of HAMA and HAMA/pdECM islets organoids. The blood vessels are stained red, and the nuclei are stained blue (scale bar = 200 µm). ‘S’ represents surrounding host tissue, ‘I’ represents the implanted hydrogel

Note. Adapted with permission from “Hyaluronic acid methacrylate/pancreatic extracellular matrix as a potential 3D printing bioink for constructing islet organoids,” by Wang D, Guo Y, Zhu J, Liu F, Xue Y, Huang Y, et al. Acta Biomater. 2022;165:86–101. (https://linkinghub.elsevier.com/retrieve/pii/S1742706122003750). © 2022 Acta Materialia Inc.

PLA

Compared with natural polymers, most synthetic polymers have super mechanical properties, and 3D printed structures can be maintained in vivo for a long time. PLA, is a widely used synthetic polymer in biomedical applications that is biocompatible, has high elasticity and mechanical strength suited for subcutaneous implantation, thus employed to promote the bio-integration of the encapsulating system [164]. Lactic acid is chiral in nature, making PLA hydrophobic and having poor cell adhesion properties. Surface treatment using plasma increases the low surface free energy of various materials and provides a solvent-free method for altering the wettability, surface chemistry and surface roughness of polymers, which enhances the cell survival and proliferation [165, 166]. The surface-free energy of PLA is also increased by plasma activation, generating a wide range of functional groups on the surface, including polar groups, which significantly alter wettability and have a favourable impact on material-cell interactions. Some studies demonstrated that plasma treatment significantly enhanced the surface’s hydrophilicity and decreased the contact angle, both of which were stable for 30 days in phosphate buffered saline (PBS) for over 30 days [167].

By optimizing the parameters of a low-cost 3D printer Song et al. developed a macroporous PLA scaffold that houses stem cell derived β-cell clusters within a degradable fibrin gel structure [168]. The diameter of cell clusters was calculated using a finite element model of cellular oxygen diffusion consumption in order to prevent severe hypoxia prior to vascularization in bioprinted constructs. Insulin was produced in response to a glucose injection following implantation of the constructs in mice, with the transplanted constructs maintaining their structural stability for 12 weeks following which they were retrievable. An important finding of the study was the non-responsiveness of the bioprinted grafts containing large cell clusters. Mice receiving implantation of these bioprinted constructs with big clusters reportedly had detectable amounts of human insulin which didn’t respond to glucose injection. Similarly, serum samples from mice receiving small clustered constructs showed the presence of human insulin. However, an approximately 2.50-fold increase in human insulin was induced by the glucose injection, showing that the transplanted grafts were functional and glucose-responsive. This strategy, as opposed to pure cell encapsulation methods, could be used as a foundation for sophisticated 3D printing strategies for diabetes cell replacement therapy.

For instance, Farina et al. [169], in their study, demonstrated that subcutaneous implantation of a functionalized, 3D-printed cell encapsulated system results in appropriate and rapid graft vascularization. PLA is used to develop an entirely novel implantable 3D printed device that will accommodate a pancreatic islets graft. Vascularization was enhanced by dispensing proangiogenic substances, such as vascular endothelial growth factor (VEGF), which is additionally known to improve islet viability and function [170]. Islets encased in this device could be protected from acute hypoxia and retain their function, resulting in insulin secretion over several months within the device. In order to meet changing physiological needs, the device’s transcutaneous refillability provides opportunities for cell augmentation without surgical extraction and re-implantation. Additionally, the reservoir structure allows for the possibility of retrieving the graft, which is crucial for designed cells made from stem cells that are experiencing malignant or other undesirable alterations.

PCL

PCL has become an increasingly popular choice among various types of biopolymers because of its multiple benefits such as biodegradability, high strength, and biocompatibility. PCL is a semi-crystalline thermoplastic polyester approved by the FDA that is manufactured by cationic or anionic ring-opening polymerization of ε-caprolactone at high temperatures [171]. Moreover, the potential for modification and functionalization with additional substrates like growth factors, and the development of a suitable scaffold design within a scalable process have convinced researchers to encapsulate cells within PCL scaffolds. Precise and complicated geometries could be created via 3D printing of PCL when additional sacrificial support materials are used. In essence, clinical islet transplantation incorporates the separation of islets from surrounding exocrine tissue in a donor via enzymatic digestion and mechanical disruption, accompanied by density centrifugation to purify the exocrine tissue [172]. Due to enzymatic degradation, pancreatic islet cells may often lose their interaction with the surrounding ECM. As a result, the natural microvasculature is disrupted causing cells to lose functionality. Although alternate transplant sites have been frequently proposed to address this challenge, the longevity of extra-hepatically transplanted islets is critically dependent on rapid re-vascularization following engraftment.

In line with a prior study by Marchioli et al. [173], developed a new 3D scaffold platform that could actively promote vascularization thereby having potential applications in extra-hepatic islet transplantation (Figure 7). Briefly, the authors prepared 3D ring-shaped PCL scaffolds surrounding an alginate hydrogel core for islet encapsulation. The heparinized surface of the PCL scaffold could bind VEGF to the alginate-encapsulated islets electrostatically, thereby improving vascularization in vivo and promoting cell adhesion because of decreased hydrophobic characteristics, greater protein binding capacity, and altered surface topography. When compared to the untreated PCL scaffold, heparin immobilization could increase VEGF retention in the scaffold by up to 3.6-fold. VEGF immobilized on the construct improved angiogenesis in close vicinity and on the surface of the scaffolds in a chicken chorioallantoic membrane (CAM) model. Moreover, islets encapsulated within the alginate core in these functionalized scaffolds showed a rounded morphology, and could be easily stained with DTZ confirming a positive response to glucose stimuli via insulin secretion, with results comparable to free-floating islets after 7 days. Overall, the aforementioned marks the development of a suitable scaffold that has the possibility of promoting rapid vascularization and endocrine function of encapsulated islets.

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Figure 7. 

3D printed PCL/alginate ring-shaped scaffolds functionalized with VEGF for islets encapsulation. (A) The hybrid scaffold concept is depicted schematically. Briefly, the islets were encapsulated in the structure’s interior using alginate hydrogel and crosslinked using CaCl₂. On the other hand, 3D plotted PCL rings adjacent to the alginate core were covalently functionalized using heparin to bind VEGF and safeguard it from degradation (left). PCL and PCL/heparin plotted scaffolds stained with Azure II (right); (B) SEM micrographs at different magnifications of the bare PCL scaffolds (top row) and of the heparinized constructs (bottom row) where the change in surface topography can be observed; (C) CAM assay using PCL and heparin-coated PCL scaffolds with three different VEGF concentrations. A 200 ng VEGF load resulted in leaking vessels on PCL scaffolds, whereas a higher VEGF concentration appeared to inhibit vessel formation. On heparin-coated scaffolds, 200 ng load VEGF appeared to induce the greatest number of blood vessels with normal morphology (arrows indicate vascularization); (D) the functional behavior of islets at day 1 (top) and day 7 (middle) when embedded within PCL/alginate hybrid scaffolds coated with heparin and functionalized with VEGF. Islets encapsulated in the alginate in the central core of the PCL/alginate scaffolds showing a round morphology and confirming the presence of insulin by staining with DTZ after 1 day of culture (bottom). SEM: scanning electron microscopy

Note. Adapted with permission from “Hybrid polycaprolactone/alginate scaffolds functionalized with VEGF to promote de novo vessel formation for the transplantation of islets of Langerhans,” by Marchioli G, Luca AD, de Koning E, Engelse M, Van Blitterswijk CA, Karperien M, et al. Adv Healthc Mater. 2016;5:1606–16 (https://onlinelibrary.wiley.com/doi/10.1002/adhm.201600058). © 2016 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim.