Integrating Additive Manufacturing (AM) Techniques to Improve Cell-Based Implants for the Treatment of Type 1 Diabetes

Abstract

The increasing global prevalence of endocrine diseases like type 1 diabetes mellitus (T1DM) elevates the need for cellular replacement approaches, which can potentially enhance therapeutic durability and outcomes. Central to any cell therapy is the design of delivery systems that support cell survival and integration. In T1DM, well-established fabrication methods have created a wide range of implants, ranging from 3D macro-scale scaffolds to nano-scale coatings. These traditional methods, however, are often challenged by their inherent limitations in reproducible and discreet fabrication, particularly when scaling to the clinic. Additive manufacturing (AM) techniques provide a means to address these challenges by delivering improved control over construct geometry and microscale component placement. While still early in development in the context of T1DM cellular transplantation, the integration of AM approaches serves to improve nutrient material transport, vascularization efficiency, and the accuracy of cell, matrix, and local therapeutic placement. This review highlights current methods in T1DM cellular transplantation and the potential of AM approaches to overcome these limitations. In addition, emerging AM technologies and their broader application to cell-based therapy are discussed.

1. Cellular Implants for the Treatment of Type 1 Diabetes

Cell-based therapies provide a potential curative approach for resolving Type I diabetes mellitus (T1DM) via the replacement of the insulin-producing cells lost to autoimmune destruction^1–3. In the native pancreas, these insulin-producing cells, termed β-cells, work to modulate metabolism and blood sugar in conjunction with other hormone-secreting cells housed within the pancreatic islet⁴. Once these cells are targeted and destroyed by the endogenous immune system during the pathology of T1DM, uncontrollable peaks in blood glucose develop, which are fatal without exogenous insulin therapy⁴. Though treatments to improve the lives of those with T1DM are evolving, most notably closed-loop systems that integrate a continuous glucose monitor (CGM) with an insulin pump, these approaches fail to replicate the complex and non-linear process of insulin secretion provided by endogenous β-cells⁵. The resulting long-term glycemic instability can lead to life-threatening health complications, including kidney disease, eye damage, and cardiovascular dysfunction⁴. To restore this metabolic control and alleviate these long-term effects, clinical islet transplantation (CIT) aims to deliver a durable treatment for T1DM through the transplantation of cadaveric donor islets^6–11. CIT protocols currently infuse these cell clusters into the hepatic portal vein, whereby islets lodge into the liver microvasculature and rapidly respond to changes in blood glucose^6–11. Systemic immunosuppression is currently the most effective method for suppressing the rejection of donor islets, but significant cellular loss associated with the transplant location, local inflammation, and insufficient immunosuppression results in sub-therapeutic dosages and generally short-term efficacy⁸.

To alleviate stressors inherent to the hepatic portal vein site, namely the physical stress of transplantation, lack of native extracellular matrix and chemical signaling, and insufficient vascularization and oxygen delivery, work has shifted to exploring alternative transplant sites and leveraging biomaterial platforms^12–19. Alternative implant sites provide different advantages, including vascular access (e.g. omental and intramuscular), accommodations for larger implants (e.g., omental or intraperitoneal), and/or less invasive implantation, retrieval, and monitoring (e.g. subcutaneous); however, no site fulfills all desired features. Biomaterial platforms can further elevate islet survival and implant safety by providing 3-D distribution and the safety of retrievability. They can also be used to locally deliver supportive agents, such as pro-survival factors to improve engraftment and/or immunotherapeutics to control inflammation and prevent immunological destruction^20–24. In addition, the polymeric encapsulation of cells within immunoisolatory hydrogels can mitigate direct immune cell attack and enhance graft protection^25,26. These biomaterial strategies for housing and supporting islets within a wide range of transplant sites have shown great promise, establishing the basis for hopeful application in humans.

The clinical translation of devices to support a therapeutic dosage of islets in humans, however, is challenged using current device fabrication methods. These methods struggle to provide adequate granularity and reproducibility. Also, the lack of control over cell, therapeutic, and overall biomaterial distribution present significant roadblocks to the creation of human-scale devices, as the variance of these features can propagate hypoxic and nutrient gradients that impact engraftment. Insufficient nutritional delivery is further exacerbated by the inability to efficiently vascularize transplanted structures and provide endogenous oxygenation and nutrients. As a result, devices are compromised by resorting to decreased thickness or cell loading and/or sacrificing a portion of the cell cluster population post-transplant²⁷. To improve factors like nutrient transport, local therapeutic release, and overall engraftment, novel fabrications methods should be implemented. Herein, we describe the emergence of a field of techniques, referred to as additive manufacturing (AM), that support the reproducible construction of complex three-dimensional structures using biocompatible materials. We detail specific AM methods, outline applications for tissue engineering, and further postulate on future applications for improving the clinical translatability and efficacy of cell-based treatments for T1DM.

2. Overview of Additive Manufacturing Methods

Additive manufacturing, or AM, is a family of fabrication techniques wherein silico modeling is used to define a complex, three-dimensional structure that is on-site fabricated through controlled methods with minimal offsite post-processing. AM takes many forms with fused filament fabrication (FFF), stereolithography (SLA), and bioprinting being the most common methods used for biomedical applications^28–31. FFF, also commonly referred to as fused deposition modeling (FDM), is a layer-by-layer fabrication method using heated thermoplastics. FFF use in rapid prototyping spans decades, as it can consistently re-create three-dimensional computer models derived from commercially available software³². More recently, this technique has found a place in medical implants through the use of biocompatible and biodegradable materials, such as poly-lactic acid (PLA), polyvinyl alcohol (PVA), polycaprolactone (PCL), or proprietary material formations³³. Compared to traditional biomaterials fabrication techniques, such as solvent-leaching and bulk-casting, FFF results in greater material geometric reproducibility and reduced variability, which are desirable features for biomedical devices. PLA and PVA FFF prints have successfully housed transplanted cells and served as rigid scaffolding for bone formation, with demonstrated long-term biocompatibility and integration^34–37. A restriction of FFF, however, is the limited resolution of the printing. Independent of the exact machining and positional systems used, the X-Y resolution of an FFF print is primarily dependent on the nozzle diameter and mechanical properties of the extruded material, resulting in minimum feature size of approximately 250 μm³³. Stereolithography, or SLA, is an alternative printing method employing light curable liquid resins, such as polyurethanes, ceramics, and hard ABS-like materials^38–40. In this approach, the resin is printed by exposing the photocurable polymer to a light source (e.g. light projector, laser) to facilitate the layer-by-layer fabrication of the desired model. Due to the higher precision provided by using light as the curing method, both the printing speed and resolution (50–150 μm, depending on the axis) are drastically increased over traditional FFF methods; however, this comes at the cost of material choice and additional processing steps for finished prints⁴¹. Biomedical applications are historically limited, as most commercial resins used in SLA are rigid and poorly evaluated for long-term soft tissue implantation. Thus, SLA has been typically used for surgical guides and dentistry applications^40–42. Imbuing light-sensitive resins with traditional biomaterials like hydroxyapatite (HAP) for bone formation has expanded their medical utility; however, heavy processing and evaluation is required to mitigate cytotoxic effects³⁹. Recently, more favorable materials such as thermoplastic polyurethane-like rubbers and custom formations have become available, permitting the fabrication of flexible scaffolds and microneedle arrays^38,43. Despite this, SLA tends to be viewed as a more specialized printing method, highly dependent on available materials and applications.

Incorporating the advantages of the core technologies from FFF and SLA, bioprinting, which is a broad term encompassing the layer-by-layer deposition of biocompatible, often biodegradable materials alongside encapsulated cells and therapeutics, seeks to provide a more versatile system. This method is attractive due to the inherent control of cell and material placement, with resolutions up to 50 μm and the potential for high cell viability (i.e., > 80%)⁴⁴. The four primary bioprinting methods include inkjet bioprinting, where droplets of biomaterials can be deposited on a substrate without physical contact using charge or other methods; laser-assisted, where irradiation of an energy-absorbing substrate can be used to deposit evaporated droplets of biomaterials onto a substrate; and more traditional methods like FDM and SLA-based bioprinting^45,46. In selecting the optimal method for a specific application, many factors come into play including desired printing speed, spatial resolution, cell viability, cell density, and cost. Although inkjet bioprinting allows low-cost printing at high speed and resolution, this method typically struggles in creating stable vertical structures due to the low viscosity of the bioinks⁴⁴. FDM and SLA-based bioprinting methods are better candidates for creating true three-dimensional structures, as their higher viscosity bioinks offer increased structural support; however, these techniques exhibit challenges, e.g. low cell viability during extrusion for FDM-based bioprinting and narrow bioink choice for SLA-based bioprinting⁴⁴. Compared to the other classic bioprinting methods, laser-assisted bioprinting methods exhibit the highest preservation of cell viability while maintaining fast printing speeds and spatial resolution⁴⁴. Although costly, laser-assisted allows for increased cell density, greater than 95% cell viability after printing, and the usage of highly viscous bioinks. As a relatively new technique, the full impacts of laser exposure on the printed cells and biomaterials within the bioinks is still being investigated, which has delayed more widespread adoption.

The application of bioprinting methods for biomedical applications is varied. For example, laser-assisted bioprinting methods are typically used for implants that require high spatial resolution, including the creation of bone, skin, and adipose tissue models^31,47. In contrast, FDM-based bioprinting is more likely to be used to create vasculature and skeletal muscle models, as cells often migrate within completed constructs during the process of self-assembling structures like vessels ^29,48. Additionally, the increased printing speeds and elevated cell viability afforded by laser-assisted and SLA-based bioprinting approaches have increased their feasibility for engineering larger tissues; however, this method is still hindered by restrictions imposed by the printer itself and currently available printing materials. Faster printing methods are needed to mitigate deleterious impacts to the cells and enhance overall scalability. Furthermore, the ability to recapitulate the in vivo environment can be improved by developing methods to employ multiple bioinks and finding ways to increase spatial resolution. Groups have explored these avenues using innovative nozzle, syringe, and motor control system techniques⁴⁹.

Overall, AM methods have a clear potential to improve the rational design and development of tissue-engineered implants. Focusing on cell-based therapy for T1DM, the need to translate device development from traditional fabrication methods to AM approaches is high, due to current roadblocks. Specifically, factors like impaired oxygen transport, lack of control over cell and material placement, and the inability to form robust and consistent vasculature continue to put restrictions on the graft size and subsequently hinder clinical translation. Applying AM methods to address these limitations could permit the fabrication of a superior insulin-secreting, cell-based implant.

3. Additive Manufacturing and Applications in β-Cell Transplantation

Optimized Implant Geometry to Improve Nutrient Transport

A primary consideration in engineering tissues for islet transplantation is maximizing nutritional transport by reducing diffusional distances between the implanted cells and the surrounding environment^15,50. While sufficient nutrient delivery is a key consideration for most cellular implants, meeting the metabolic demands for islets, particularly concerning oxygen, is arduous⁵¹. Previous work has focused on the role that the parameters of geometry and pore size play in improving the mass transport of critical metabolites such as oxygen and glucose. Traditional scaffold fabrication methods introduce macro-scale porosity through techniques like porogen leaching, electrospinning, and gas foaming⁵². Resulting highly porous (> 85% v/v) materials provide physical support and distribution of the islets, while also increasing nutritional access by both the passive transport of oxygen through their inherent open space and the facilitation of host intra-device vascularization^53–55. Modulating pore size using these methods is feasible via modification of porogen size and loading, which alters both the overall pore size and pore interconnectivity^56–59. In addition to improved nutrient transport, implant porosity globally impacts biocompatibility and engraftment long-term, with optimal pore sizes facilitating healthy cellular infiltration and remodeling^20,21. The intimate control of pore size and connectivity using particulate leaching and gas foaming methods, however, is limited, resulting in heterogeneity of the final scaffold product. In addition, uneven geometry as a result of the irregular porogen surface can lead to rough macro-scale topographies, potentially increasing the risk of proinflammatory responses^60,61. To enhance control over both macro and microporosity, an alternative approach combined electrospinning with photoresist molds and laser drilling to achieve both high porosity and a niche-based cell spheroid distribution ^62,63. The resulting devices prevented islet aggregation and retained islet function. While this more complex method elevated the reproducibility of porosity and islet distribution, the fabrication of new molds remains time-consuming and non-modular.

Alternative to open scaffolds, the encapsulation of cells within nano- or micro-porous hydrogels that impair the direct interactions of host and implanted cells is a common transplant technique for insulin-producing cells, as it can mitigate the immunological rejection of the implant^64–69. Encapsulation can range from nano-scale coatings to macroscale hydrogels and can be fabricated from materials such as alginate, poly(ethylene glycol), and agarose⁶⁴. Nanoencapsulation or ultrathin coatings (<20 μm) can permit the protection of cells; however, incomplete encapsulation and complexity of fabrication can make these methods difficult to scale^67,70. At the microencapsulation scale, islets can be housed in hydrogels 400–900 μm in diameter, which delivers more complete encapsulation at the cost of hindered mass transport to the encapsulated cells.^64,65 At the macroscale, fabrication methods are less arduous and the retrieval of transplanted cells in the event of graft failure becomes more feasible but at the cost of significantly delayed nutritional diffusion through the thicker hydrogels, resulting in hypoxic gradients.^64,65,71 Geometric manipulation to increase surface area and global mass transport efficiency is one approach to improve transport. For example, Wang et. al. created a single micron-scale encapsulation tube for housing insulin-producing clusters by augmenting alginate with electrospun nanofibers, resulting in an efficacious implant.⁷² Alternative methods that leverage additional technologies to improve nutrient delivery, such as the integration of oxygen-generating materials and/or pro-vasculogenic factors, can improve oxygen access for the cells within traditional macroencapsulation devices^73,74. However, the integration of these features discreetly and consistently for rapid fabrication using traditional methods is challenging.

AM approaches can support the integration of more complex, but reproducible, geometries and features to mitigate mass transport roadblocks. At a surface level, AM allows for the fabrication of scaffolds from historically validated biomaterials but with defined and consistent geometry based on distinct design parameters. This robust control delivers a defined macro-scale geometry and micro-scale pore size, allowing the user to tailor their materials to different applications^75–77. From a purely engineering perspective, simply improving the total surface area to volume ratio for an implant design should drastically improve oxygen and nutrient diffusion. For example, to modulate the global construct shape and provide consistency of the macro-implant, Grattoni’s team used a traditional FFF method to print a PLA-based implant with discreet spatial patterns at high resolution^78,79. The internal 300 μm pore size scaffolds provided defined spaces for distributing individual islets within the device inner chamber, while smaller external micro-pores supported subcutaneous perfusion and dedicated inputs for islets infusion⁷⁸. Additional studies using the scaffold to house testosterone-secreting Leydig cells, in combination with a VEGF-doped platelet-rich plasma hydrogel, reported successful device vascularization and sustained cellular survival in a murine, subcutaneous transplant model. Ernst et. al. further explored this concept for macroencapsulation implants by applying a hydrogel coating onto a toroidal AM-elastomer backbone³⁸. This interlocking, toroid shape, printed using a flexible proprietary resin, increased the macroscale total surface area, thereby permitting more efficient oxygen delivery to the encapsulated cells due to the decreased diffusional distance. Compared to control spherical hydrogels, this new geometry allowed for improved mass transport, which yielded increased cell viability post-encapsulation in the alginate hydrogel. Implantation of these connected rings restored normoglycemia in diabetic murine recipients. Multiple studies have solidly demonstrated the expected benefit of AM in improving the flexibility in device geometry for improved mass transport^80,81.

Improved Vascularization and Engraftment

Rich host vascularization of devices is necessary for the long-term survival of most transplanted cells. For β-cell therapy in T1DM, the competency and consistency of the vascular network are essential to not only support efficient nutrient delivery but to ensure effective insulin kinetics. In earlier platforms, the infiltration by endogenous cells and subsequent anastomoses with islets was directed by modifying bulk material properties to create pore and/orsurface geometric features that encourage vessel formation^{20,53,82–86} These approaches were successful in generating robust vasculature; however, islet loss post-transplantation during the delay in the formation of competent vascularization remains a challenge. In response, work has shifted to “priming” the transplant site by pre-vascularizing biomaterials with endogenous cells or generating vessels in vitro for accelerated vessel development^87–89. These approaches have significantly improved the efficiency in the development of a competent vascular bed. Continuous improvements using new hydrogel formulations, novel therapeutics, and culture techniques, like co-cultures, have further advanced the field^14,15,90,91. Leveraging advanced manufacturing methods, however, can further optimize micro-scale geometry to not only accelerate graft anastomosis with host vasculature but control intra-device vascular structure and homogeneity.

The design of vascular structures using AM methods is a long-standing interest in tissue engineering device development. Initial work using FDM printing methods focused on reproducing macro-scale vascular structures. For example, Visser and colleagues created large-scale vessel-like conduits with macroscale branching structures using multi-material FDM printing of PVA and PCL as the support and printing material, respectively⁴⁹. Using an SLA-based printing method, Meyer et. al. fabricated flexible, tubular structures made from proprietary materials⁹². Though these structures could be used to replace large sections of damaged tissue, their scale was not highly applicable to cell-based transplants. In response, groups such as Grattoni and Niklason leveraged traditional FDM printing techniques to create patterned structures to guide vasculature formation both in vitro and in vivo^78,79,86. This process permitted Niklason’s group to interrogate the effect of pore size on hSMC vessel formation and Grattoni’s group to create a subcutaneous device for transplantation and subsequent vascularization of human islets in a nude mouse model. AM-based fabrication resulted in improved implants, as increased printing resolution allowed for the control of microscale pore size and the introduction of more bio-active/degradable materials such as poly(lactic-co-glycolic acids) (PLGA). These systems supported rapid development and testing; however, the use of more rigid materials is not optimal for long-term soft tissue implantation. Shifting to bioprinting technologies, complex and branching conduits can be formed using supportive hydrogels, such as collagen or methacrylate gels, which allow cells to sprout throughout the matrix under the desired geometric constraints^93–96. These approaches not only support the investigation of the biology of vessel formation, but they lay the groundwork for the eventual clinical translation of these techniques to a broad swath of cell-based therapies. Islet-specific work in the AM-vascularization space has primarily focused on fundamental biological discoveries, with groups like Hospodiuk et. al. using FDM technology to explore vessel sprouting from pseudo islets ⁹⁷. While publications creating macroscale, AM-manufactured vascularized constructs for islet transplantation are currently limited, advancements made in leveraging AM to support favorable vascularization conditions, such as the selection of optimal pore sizes and optimal material choices, facilitate their translation to β-cell therapy in T1DM.

Controlled Cell and Matrix Placement

Beyond the utilization of AM-based methods to improve three-dimensional biomaterial distribution, there is potential for micro-level control of cell and protein distribution. This degree of control would benefit the design of T1DM implants containing β-cells, as their cellular function is highly dependent on factors like oxygen access and interactions with critical extracellular matrix (ECM)-binding proteins^98–100. Providing optimal cellular distribution, as well as a supportive ECM niche, would reduce the rapid loss of function and anoikis-driven apoptosis commonly observed post-transplantation. Most traditional scaffold fabrication methods load the cells post-device fabrication, which leads to broad heterogeneity of cellular distribution and challenges in loading larger cellular spheroids. Furthermore, separating the material fabrication and cell loading process leads to elevated bulk of the final implant volume. Traditional encapsulation methods that support in situ material and cell placement reduce these issues but struggle with unpredictable cellular distribution within the material and incomplete encapsulation. These issues are exacerbated when incorporating large cellular spheroids, such as islets. For example, alginate islet microbead encapsulation using fluidic or electrostatic bead generation methods commonly results in either heterogeneous clusters per capsule for large microbead sizes or incomplete encapsulation for smaller microbead sizes.¹⁰¹ Seeking to support the loaded cells through the integration of ECM components further complicates device fabrication. While numerous publications have shown improved islet function when incorporating ECM proteins such as laminin, collagen, and fibronectin, most studies present these components through surface coatings of traditional biomaterials or bulk hydrogels, limiting their homogenous presentation across the graft.^102–104

AM-based bioprinting approaches provide an avenue to remove these limitations. Specifically, bioprinting supports the in situ encapsulation and printing of multiple cell types and ECM while simultaneously controlling 3D placement and dimensions. This utility was clearly demonstrated by works such as that of Pati et. al. who printed decellularized ECM (dECM) hydrogels alongside cells, forming a tissue analogue with different properties¹⁰⁵. In this setup, cells were exposed to different proteins depending on the desired geometry and dECM distribution, allowing for the enhanced expression of targeted phenotypic genes, e.g. cardiogenic, adipogenic, and chondrogenic, when compared to just collagen matrix. Preliminary work additionally demonstrated the ability to print multiple dECM bioinks into a single construct, opening the door for more complex tissue analogues containing discreet spatial ECM patterns. The work of Kim et. al. further demonstrated the benefit of spatial control through the fabrication of a skin patch with a defined dermis and epidermis region, which was printed using discreet cell types and dECM¹⁰⁶. Though in its relative infancy, this methodology is also moving into the diabetes space with groups like Liu et. al. developing printing platforms suitable for islets¹⁰⁷. Using a custom-made coaxial printer, they created a core-shell structure where islets could be printed in alginate, gelatin, and methacryloyl hydrogels that were surrounded by endothelial progenitors. This resulting device delivered an immunoisolatory platform of defined geometrical scales containing uniformly distributed islets as well as supportive cells that could theoretically promote peripheral device vascularization once fully engrafted. Going further, there is potential for non-endogenous bioactive materials to be printed alongside islets as well. Examples like Hu et. al. highlight the ability to print β-cells into hydrogels doped with immune-regulatory components such as pectin¹⁰⁸. Though studies involving islets are still early, advancements in bioprinting of cell-based constructs show the potential of this approach to improve the physical and chemical environment for transplanted cells through controlled matrix protein and cell placement. Such approaches would also elevate the consistency of the therapeutic dosage, as well as the final product.

Controlled Integration of Therapeutic Agents

Parallel to cell and matrix placement in cell-based construct design, the notion of using local therapeutics to regulate cell behavior and enhance transplantation platforms has grown in popularity in recent years. Strategically placing therapeutics into these islet implants has the potential to help recapitulate the in vivo environment by assisting in multiple critical roles, including angiogenesis, pro-islet health, and immunomodulation. Angiogenesis has long been a primary target for cell-based therapies due to the inherent requirement for the transport of oxygen and nutrients to the graft. Facilitation of vessel development usually employs a critical growth factor, for example, vascular endothelial growth factor (VEGF), to accelerate the formation of endogenous vessels²². In addition, the incorporation of pro-islet therapeutics, such as small molecules (e.g., curcumin, exenatide, and vitamin D) and hormonal therapeutics (e.g. 17β-estradiol), can lead to decreased apoptosis and elevated and/or more durable insulin secretion^20,109–112. Finally, local immunomodulation broadly involves the local release or presentation of agents that dampen the innate or adaptive immune response to transplanted cells, which would support enhanced engraftment and long-term survival while reducing the need for systemic immune suppression. These therapeutics can take many forms, ranging from small-molecule steroids (dexamethasone, fingolimod)^{20,111,113,114}, to larger macrolides (rapamycin, tacrolimus)^115,116 and proteins (PD-L1)¹¹⁷. Overall, these therapeutics can provide significant improvements when delivered in the context of local release or presentation; however, their delivery is typically limited to homogenous distribution within the graft. This limits spatial customization of the graft site and can lead to deleterious effects on the transplanted cells or the host response. Specifically, numerous immunomodulating therapeutics impart negative effects on islet health when delivered in close proximity, therefore proper consideration of dosage is crucial for success^21,113. Furthermore, the local delivery of potent anti-inflammatory agents can lead to suppression of implant engraftment, as the host cells are impaired in migrating into the implant^20,21. Finally, the spatial control over the delivery of therapeutic agents provides a unique means to create discreet gradients within and around the implant site, thereby delivering more nuanced and physiological cues to the cells of interest.

The spatial control over therapeutic delivery concerning transplanted cells could provide new avenues for the discreet modulation of both host and transplanted cell responses. AM methods, particularly bioprinting techniques, are a likely candidate for solving these issues. For example, Tarafder et. al. combined traditional depot-based drug release with bioprinting to achieve distinct control over therapeutic location¹¹⁸. This melding of PLGA microspheres with PCL allowed for the spatial release of transforming growth factor beta-3 (TGF-β3) and connective tissue growth factor (CTGF), leading to the generation of heterogeneous in vitro tissues. Similarly, Liu et. al. combined bioprinting with electrospinning to create scaffolds with simultaneous release of bone morphogenetic protein-9 (BMP-9) and VEGF for osteochondral tissue generation¹¹⁹. Alternatively, Freeman et. al. fabricated composite hydrogels capable of simultaneously releasing VEGF and BMP-9 with a construct-wide gradient¹²⁰. This enabled them to experiment with different loadings to find the optimal setup for in vivo angiogenesis. Though these methods have yet to be applied directly to islet transplantation, their use in conjunction with previously mentioned techniques shows promise for their translation. Specifically, the localization of a potent modulating therapeutic to the periphery of a device would allow for enhanced targeting of the host immune system while limiting the agent’s potentially negative effects on islets. In addition, pro-islet therapeutics could be simultaneously loaded within the matrix space more closely localized with the islets, which could enhance their impact. This prospective application, therefore, has untapped potential for improving the long-term outlook of cell-based devices for the sustained treatment of T1DM.

4. Emerging Technologies and Clinical Aspirations

As an emerging field, there continues to be iterative improvements to AM methods that enhance accessibility and decrease costs; however, inherent limitations on the scale of fabrication and material choice are remaining roadblocks inhibiting their eventual translation to tissue-engineered clinical products. To address these obstacles, unique AM methods have emerged in the healthcare space that present novel solutions to these problems.

Increased Scale of Bioactive Constructs

A crucial limitation preventing the translation of bioprinting to the human scale is the inability to print large, complex constructs using hydrogel-based materials. In typical bioprinting, hydrogels are extruded as a semi-viscous liquid then cross-linked into its printed orientation to form three-dimensional structures. Thermal or UV crosslinking is often used due to its greater compatibility with encapsulated cells; however, in the time from extrusion to full crosslinking, gravity pulls and distorts the structure, making the fabrication of large constructs with overhangs impossible. As a result, groups often default to a single pattern, commonly termed the ‘crosshatch’, to make their constructs^{35,76,80,120–122}. Recent advancements in the printing medium, however, have addressed this challenge. One example of interest is the work of the Feinberg group using supportive matrices ¹²³. Described as the freeform reversible embedding of suspended hydrogels, or FRESH, large-scale printing of soft bio-inks is made possible within a supportive gelatin matrix. Once the bio-ink is cross-linked, the support matrix can be non-destructively melted away to release the finished structure. This methodology allows for traditionally impossible structures with complex overhangs and void spaces to be printed, permitting the fabrication of structures on the scale of a human heart using biocompatible polymers^123,124. Another approach in 3D soft material printing has emerged from the Angelini group. By printing within a matrix of small hydrogel spheres as supports, termed jammed microgels, semi-viscous materials can be printed in three-dimensional space. This enables the high accuracy printing of complex models using biocompatible inks like collagen, as well as stable elastomer structures such as organosilicones^125,126. With the development of these new printing approaches, challenges with bioprinting scalability are largely being addressed.

Optimized Multi-material Printing

As groups strive to further recapitulate the native microenvironment of transplanted cells to improve engraftment, the printing of multiple biocompatible materials has gained popularity. As discussed previously, highly complex cell-based therapies are likely to require multiple cells/cell cluster types, therapeutics, and materials to achieve the desired in vivo effect, so the implementation of streamlined fabrication techniques will be essential. A potential solution to this is single-tool, multi-material fabrication. Typically, FDM-based fabrication methods (FFF/FDM, FDM-bioprinting) utilize multiple tools to print two or more different materials/inks for a single device. This requirement results in greater complexity and printing time to achieve seamless results while also increasing the cost of entry. To remove these penalties on multi-material prints, groups like Skylar-Scott et. al. have implemented single nozzle solutions¹²⁷. Using a combination of gas-based extrusion and a custom-made manifold, flexible silicones can be extruded discretely or mixed within a single tool head to yield complex prints at accelerated rates. Moving to more biocompatible materials, Liu et. al. utilized a similar multi-material manifold to extrude bioinks¹²⁸. Specifically, they fabricated bulk hydrogels with up to four discrete “zones” of cells, though the height of their structures was limited due to the lack of support material. As these methods of fabrication become more optimized, opportunities will arise to significantly improve the printing of more sensitive cell types, such as islets, by minimizing the printing time and better tailoring the surrounding microenvironment.

Current Pathways and Roadblocks to Clinical Applications

AM-based approaches, to date, have demonstrated significant potential to deliver superior device designs for a broad spectrum of tissue engineered approaches. There is now a need to push these AM devices to the clinic. Groups such as Paez-Mayorga et. al. and Wang et. al. are contributing to this translation by testing AM fabricated devices within more complex and larger scale preclinical models, such as non-human primates^129–131. Product development and clinical application of AM-based devices are slowly increased in prevalence, although most devices are used as surgical guides with only the recent translation of an implantable acellular graft for hard tissue growth^132,133. Concurrently, regulatory guidance for AM fabricated acellular or combinatory tissue-engineered products is evolving in the US, EU, and Japan^134,135. For example, the FDA recently released a discussion paper on the potential benefits and risks associated with the fabrication of devices on-site for implantation¹³⁶. It is anticipated that the AM features of enhanced reproducibility and control may lead to reduced challenges in regulatory approval; however, leveraging the unique capacity of AM-based approaches to create devices with highly complex features may alternatively result in a more complicated approval process.

While there are no current AM-based devices in clinical trials for T1DM cell therapy, recent trials using stem cell-derived β-cells, specifically those from Vertex Pharmaceuticals (ClinicalTrials.gov identifier: NCT04786262) and Viacyte (ClinicalTrials.gov identifier: NCT03162926, NCT04678557), have demonstrated both market interest and therapeutic potential in this area. Although current results implicate lack of full resolution of T1DM in patients due to sub-therapeutic dosing, hallmark metrics such as elevated serum c-peptide and decreased exogenous insulin dose indicate the feasibility of a theoretically unlimited source for T1DM cell therapy; a significant advancement that would allow for wide scale use^137–139. Unfortunately, current stem cell-based implants are still subject to the same major implantation roadblocks experienced by previous CIT implementations, notably inadequate physical and chemical stimuli, improper or delayed vascularization, and the requirement for constant systemic immunosuppression to prevent graft loss. As highlighted herein, AM-based methods have an opportunity to be disruptive by delivering enhanced geometric modularity and complexity of the final device design. Simultaneously, AM-based control over cell and matrix placement can serve to both enhance cell survival and provide a more reproducible final combinatory product. Finally, innovative immunoprotective features can be integrated into the design, including controlled material encapsulation and placement of local drug delivery depots. Future work should seek to translate the utility of AM methods to improving the safety, reproducibility, and functionality of stem cell-based implants for T1DM.

The enhanced 3D design features AM provides theoretically supports the customization of device parameters on a patient-by-patient basis, a trait that could prove invaluable as the field moves towards more personalized medicine and practices. This modularity, however, will likely complicate the regulatory process. Modifying parameters inherently introduces variability, which exponentially grow with each additional parameter. This will likely make standardizing device release criteria for regulatory approval more difficult than one-size-fits-all approaches. To accelerate clinical translation, emerging AM-based cell therapy devices will likely need to focus on single design prototypes prior to moving to customized products. As an additional challenge, emerging T1DM cellular platforms are often combinatory products that use not only multiple biomaterials but also several therapeutics, cell types, and proprietary printing setups. Thus, streamlining the regulatory process for these combinatory AM-based products will likely be the largest hurdles in eventual clinical application.

6. Conclusions

While AM adoption in the translational space is in the early stages, there is clear potential for disruptive discoveries and designs to improve on traditional cell-based transplant platforms for the treatment of T1DM. Due to the highly tunable nature of technologies like FDM, SLA, and bioprinting, there is exponentially greater flexibility for not only initial design and fabrication but also subsequent optimization of construct conditions. Cell-based therapies, especially islet transplantation, have faced fundamental roadblocks to widespread clinical adoption due to factors like hypoxia, immune rejection, and constrained transplant size. With new AM tools available that deliver enhanced control over all aspects of implant design, there is potential for new ideas in the CIT space that can substantially enhance the clinical efficacy and reproducibility of these implants.

Figure 1.

Overview of AM-based methods for fabrication of cell-based transplant platforms

Figure 2.

Improvements to cell-based platforms provided by AM-based approaches focused on elevating nutrient mass transport and graft vascularization

Figure 3.

Applications of AM methods for improved control over cell, matrix protein, and therapeutic placement