Angioplasty devices and methods for transplant diagnostics and therapy - Patents.com

Abstract

The present disclosure describes new devices and methods for angiogenic devices and methods for implanted diagnostics and therapeutics. In certain embodiments, the present disclosure provides devices comprising a degradable shell and a non-degradable core. In certain embodiments, the non-degradable core can contain encapsulated therapeutic cells and / or biosensors, and the degradable shell serves as a scaffold for vascular growth and enhances blood flow to the cells and / or biosensors.

Description

[Technical field]

[0001] explanation Claiming priority This application claims the benefit of priority to U.S. Provisional Application Serial No. 62 / 806,496, filed February 15, 2019, the entire contents of which are incorporated herein by reference.

[0002] Government Assistance Clause This invention was made with Government support under Grant No. R01DK120459 awarded by the National Institutes of Health. The Government has certain rights in this invention.

[0003] A. Field This disclosure relates to the fields of biology, medicine, medical devices and immunology. More specifically, it relates to angiogenesis devices and methods for their use as implant diagnostics and therapeutics. [Background technology]

[0004] B. Related Technologies The field of tissue engineering aims to meet the great clinical need for the development of techniques and processes that promote regeneration from human disease or injury by incorporating a combination of cells, biomaterials, and biochemical and physicochemical factors. Indeed, a tremendous amount of research has been done over the past decades to construct implantable vascularized constructs such as skin, cornea, and bladder. However, obtaining functional and physiologically relevant tissues remains a major challenge in the field, as the vasculature is required to supply nutrients and remove waste products in thick constructs. This challenge can be attributed to limitations in diffusive transport, leading to necrotic cells due to insufficient access to nutrients. Furthermore, the lack of ability to reproduce heterogeneous patterns of cells and matrices to obtain constructs with controlled geometry and architecture has hindered progress towards organ replacement.

[0005] Several approaches have been investigated to facilitate the transport of nutrients into engineered tissues. One of the most common approaches utilizes various material processing steps (such as critical point drying, gas foaming and salt leaching, or electrospinning) to form macroporous structures that can be perfused in vitro for tissue culture. Although these processes match the mechanical compliance of native tissues and provide templates for constructing functional tissues with high surface areas for nutrient and oxygen delivery, precise spatial control of scaffold architecture is not easily achieved with these methodologies, given that these constructs contain open, porous void volumes rather than vascular networks. Moreover, most of the processing steps involved during the fabrication of these engineered tissues require cytotoxic reagents or conditions, further limiting their use in the presence of living cells.

[0006] Additionally, needle-based modeling techniques have been used to create straight, perfusable channels in natural and synthetic materials. However, due to the use of straight needles, this approach does not allow for sophisticated constructs that closely resemble the heterogeneity and complexity of tissue.

[0007] Another approach utilizes soft lithography, a set of techniques adopted from the microprocessor industry. It offers spatial control of scaffold architecture through photolithographic light patterning, resulting in microfluidic networks. However, these techniques require expensive proprietary equipment to reach micrometer-scale resolution, and this mode of fabrication is too slow to efficiently fabricate large models of organ vasculature. Furthermore, lithography is typically performed in Cartesian coordinates, resulting in channels with rectangular cross-sections. However, native vasculature rarely follows linear x, y, or z vectors, and blood vessels have circular cross-sections. To obtain multilayered engineered tissues, strategies that combine photolithographic processes with layer-by-layer lamination methods have been pursued. Unfortunately, the harsh fabrication process and the difficulty of aligning successive layers with high precision, as well as the presence of edge-to-edge artifacts, limit the use of this approach to obtain complex three-dimensional (3D) interconnected microfluidic networks.

[0008] To circumvent the challenges associated with manual alignment of layers, multilayered 3D structures have been fabricated in a layer-by-layer fashion, either manually or automatically, using 3D projection stereolithography to fabricate an entire layer in a single exposure (Lin et al., 2013) or deposition techniques to fabricate heterogeneous 3D tissue constructs (Kolesky et al., 2014). Although these efforts have yielded engineered tissues with spatially patterned cells (Gauvin et al., 2012), they do not resemble the complex, physiologically relevant vascular structures. Moreover, so far, only a few layers of hydrogels have been fabricated.

[0009] Although the biological vasculature consists of fluid networks ranging in size from a few centimeters (e.g., the aorta) to a few microns (e.g., capillaries), currently available techniques cannot fully capture the full range of transport networks important for multicellular organisms. Furthermore, most current in vitro models contain 3D constructs with a single vascular network containing one inlet and one outlet. However, the presence of interpenetrating vascular networks, a key feature of complex biological systems, is central to mammalian physiology for sufficient transport of nutrients, although it is absent from the current literature. For example, alveoli must be located in close proximity to the pulmonary capillary network to ensure oxygenation of red blood cells due to the nature of passive transport of oxygen.

[0010] Thus, a fundamental barrier to successful cell therapy is the lack of biocompatible implantation devices. Implanted biomaterials trigger a foreign body response (FBR), a cascade of inflammatory events and wound healing processes that lead to fibrosis (compartment formation) and subsequent implant failure (Harding and Reynolds, 2014; Langer, 2009; Williams, 2008). This response has been described for many materials, from naturally occurring polymers to synthetic materials (Ratner, 2002; Ward, 2008; Zhang et al., 2013). Thus, there is a significant medical need to develop biomaterials that overcome this key challenge of eliminating the FBR associated with biomaterial implantation. We have addressed this challenge through the identification of immunomodulatory small molecules by combinatorial chemistry and high-throughput screening (Vegas et al., 2016b) and have validated the long-term efficacy of these small molecules in preventing fibrosis and long-term cell survival in non-human primates.

[0011] The most commonly studied method for islet encapsulation therapy is the formulation of isolated islets into alginate microspheres (Lum et al., 1991; Scharp and Marchetti, 2014; Schneider et al., 2005). Alginate is a dicationic (Ca 2+ , B.A. 2+ Alginate microspheres are versatile biomaterials that form hydrogels in aqueous solutions and are used in several biomedical applications such as drug delivery, tissue regeneration, implantable sensors, and encapsulation (Kearney and Mooney, 2013, Lee and Mooney, 2012). The popularity of this coating and encapsulation technique lies in its low cost, low toxicity, gentle gelation that is harmless to cells, and persistence. Unfortunately, immune recognition of even empty alginate microspheres can elicit FBR, which can be further exacerbated by the presence of encapsulated allogeneic or xenogeneic donor tissue (de Vos et al., 2006, Robitaille et al., 2005, King et al., 2001). This fibrotic response to alginate has been observed in experimental studies in C57BL / 6 mice and non-human primates (NHPs) ( Jacobs-Tulleneers-Thevissen et al., 2013 , Scharp and Marchetti, 2014 , King et al., 2001 ), as well as in human clinical trials.

[0012] Furthermore, the requirement to place allogeneic islet cells near the bloodstream to rapidly respond to blood glucose levels, especially during insulin secretion, coupled with the high oxygen demands of islet cells, adds further challenges to recapitulating normal pancreatic function on a whole-organ scale. Although these angiogenic challenges are well defined, biomaterial technologies that could be used to develop angiogenic devices remain unclear. Summary of the Invention

[0013] overview Briefly, the present disclosure provides angiogenic devices and methods for transplant diagnostics and therapy. We applied modern 3D bioprinting technology to induce angiogenesis of immunoisolated therapeutic cell constructs for long-term survival and functionality, allowing for higher density engraftment of therapeutic cells.

[0014] In light of the requirements for an effective vascular network, design criteria for fabricated tissue that mimics vascular geometry include, but are not limited to, multiple branching vessels with sizes ranging from tens to hundreds of micrometers in diameter, and substantially smooth inner walls that can minimize frictional resistance and turbulence during fluid flow.

[0015] The present disclosure provides, in certain embodiments, devices comprising a degradable shell and a non-degradable core, in which the non-degradable core can contain encapsulated therapeutic cells and / or biosensors, and the degradable shell acts as a scaffold for vascular growth to enhance blood flow to the cells and / or biosensors.

[0016] More specifically, an implantable device is provided that includes a degradable outer housing that includes a degradable polymer and an inner non-degradable portion that includes a degradable polymer. The degradable outer housing can include a patterned architecture that promotes vascularization of the implantable device (i.e., angiogenic bed) and / or the non-degradable portion can include a natural or synthetic hydrogel. The synthetic hydrogel can be poly(ethylene glycol) diacrylate (PEDGA). The device can further include gel methacrylate (GelMA), natural and chemically modified hyaluronic acid, and natural and chemically modified alginate.

[0017] The hydrogel may contain more than 80% water. The hydrogel may contain about 10% GelMA and about 3% PEDGA. The hydrogel may contain acrylate-(PEG-protein) n-PEG-acrylate monomers, where optionally PEG is 2-35 kDa and the protein is greater than 100 Da and the total molecular weight is at least 500 kDa or n=2-500 repeat units. The hydrogel may comprise acrylamide, vinyl sulfone, norbornene substituted acrylate groups, alginate, silk, dextran, chondroitin sulfate, hyaluronic acid, cellulose, heparin, poly(caprolactone) or multicomponent versions thereof.

[0018] The inner non-degradable portion may be loaded with drugs, i.e. small molecule therapeutics or biological therapeutics (peptides, antibodies, etc.), crystals, cells (cells from xenogeneic tissue, cells from cadavers, stem cells, cells derived from stem cells, cells from cell lines, primary cells, reprogrammed cells, reprogrammed stem cells, cells derived from reprogrammed stem cells, genetically engineered cells, etc.), or combinations thereof. The cells may be human cells and / or insulin-producing cells and / or pancreatic islet cells. The inner non-degradable portion may be patterned. The device may include multiple layers of hydrogel.

[0019] The degradable outer housing or the inner non-degradable portion can be loaded with drugs, i.e., small molecule therapeutics or biological therapeutics (peptides, antibodies, etc.), crystals, or oxygen-releasing compounds such as CaO2. The degradable outer housing or the inner non-degradable portion is loaded with non-crystallized, crystallized, and / or encapsulated pro-angiogenic factors (such as VEGF), or pro-angiogenic cells (endothelial cells, mesenchymal stem cells, endothelial progenitor cells, pro-angiogenic macrophages, etc.). The inner non-degradable portion may be loaded with biosensors, such as biosensors capable of measuring blood oxygen, blood glucose, tumor antigens, cholesterol, creatinine, hemoglobin A1C, insulin, glucagon, hormone levels, levels of small molecule drugs, DNA, RNA, cytokines, circulating tumor cells, among other possible diagnostic indicators. The non-degradable outer housing and / or the inner non-degradable portion may be patterned (such as where the inner non-degradable portion is patterned to allow for angiogenesis once implanted), and the patterned architecture may include vascular geometries such as multiple branching vessels ranging in size from tens to hundreds of micrometers in diameter and may have substantially smooth inner walls to minimize frictional resistance and turbulence as fluids flow through them.

[0020] Also provided are methods of monitoring a state or condition of a subject, comprising implanting a device described herein into the subject, the device comprising a sensor or biosensor; methods of treating a medical condition or disease in a subject, comprising implanting a device described herein into the subject, the device comprising a therapeutic agent; methods of creating an angiogenic device in a subject, comprising implanting a device described herein into the subject; methods of creating angiogenic tissue around an implant or prosthesis in a subject, comprising implanting a device described herein into the subject, the implant or prosthesis being such as a neural interface device; and methods of creating augmented tissue around an implant or prosthesis in a subject, comprising implanting a device described herein into the subject, the implant or prosthesis being such as a breast implant or tissue filler.

[0021] Any embodiment of any of the present devices, methods, compositions, kits, and systems can consist of or consist essentially of the recited steps and / or features, rather than comprise / include / contain / have. Thus, in any of the claims, the terms "consist of" or "consist essentially of" can replace any of the above open-ended linking verbs in order to modify a given claim from the extent to which an open-ended linking verb would otherwise be used.

[0022] Use of the term "or" in the claims is used to mean "and / or" unless expressly indicated to refer to alternatives only or where the alternatives are mutually exclusive, however, the present disclosure also supports a definition that refers to alternatives only and "and / or". Throughout this application, the terms "about" or "approximately" are used to indicate that a value includes the standard deviation of error for the device or method being used to determine the value. In accordance with long-standing patent law, the words "a" and "an" when used in conjunction with the word "comprising" in a claim or specification indicate one or more unless otherwise indicated.

[0023] [The present invention 1001] a degradable outer housing comprising a degradable polymer and an inner non-degradable portion comprising a non-degradable polymer; 4. An implantable device comprising: [The present invention 1002] The implantable device of the present invention 1001, wherein said degradable outer housing comprises a patterned architecture that promotes vascularization of said implantable device. [The present invention 1003] The implantable device of claim 1001 or 1002, wherein the non-degradable portion comprises a natural hydrogel or a synthetic hydrogel. [The present invention 1004] An implantable device according to claim 1003, wherein said synthetic hydrogel is poly(ethylene glycol) diacrylate (PEDGA). [The present invention 1005] The implantable device of the present invention 1003 or 1004, further comprising gel methacrylate (GelMA). [The present invention 1006] An implantable device according to any one of claims 1003 to 1005, wherein the hydrogel contains more than 80% water. [The present invention 1007] The implantable device of any of claims 1004 to 1006, wherein the hydrogel comprises about 10% GelMA and about 3-4% PEDGA having a molecular weight of about 3.4 kDa. [The present invention 1008] The hydrogel is an acrylate-(PEG-protein) n -An implantable device according to any of claims 1001 to 1003, comprising a monomer of PEG-acrylate, optionally wherein the PEG is 2 to 35 kDa and the protein is greater than 100 Da, with a total MW of at least 500 kDa or n = 2 to 500 repeat units. [The present invention 1009] The implantable device of any of claims 1001 to 1003, wherein the hydrogel comprises acrylamide, vinyl sulfone, norbornene-substituted acrylate groups, alginate, silk, dextran, chondroitin sulfate, hyaluronic acid, cellulose, heparin, poly(caprolactone), or multi-component versions thereof. [The present invention 1010] An implantable device according to any one of claims 1001 to 1003, wherein the inner non-degradable portion is loaded with cells, drugs (i.e., small molecule therapeutic agents or biological therapeutic agents (peptides, antibodies, etc.)), or crystals. [The present invention 1011] The implantable device of the present invention 1010, wherein the cells comprise cells from heterologous tissue, cells from a cadaver, stem cells, cells derived from stem cells, cells from a cell line, primary cells, reprogrammed cells, reprogrammed stem cells, cells derived from reprogrammed stem cells, genetically engineered cells, or combinations thereof. [The present invention 1012] The implantable device of the present invention 1010, wherein the cells are human cells. [The present invention 1013] The implantable device of the present invention 1010, wherein the cells are insulin-producing cells. [The present invention 1014] The implantable device of the present invention 1010, wherein the cells are pancreatic islet cells. [The present invention 1015] The degradable outer housing or the inner non-degradable portion is preferably made of CaO 2 The implantable device of any one of claims 1001 to 1014, which is loaded with an oxygen-releasing compound such as [The present invention 1016] The implantable device of any of the present inventions 1001 to 1015, wherein the degradable outer housing or the inner non-degradable portion is loaded with a drug (i.e., a small molecule therapeutic agent or a biological therapeutic agent (peptide, antibody, etc.)), a crystal (having non-crystallized, crystallized, and / or encapsulated pro-angiogenic factors, such as VEGF), or a cell such as a pro-angiogenic cell (endothelial cell, mesenchymal stem cell, endothelial progenitor cell, pro-angiogenic macrophage, etc.). [The present invention 1017] The implantable device of any one of claims 1001 to 1009, wherein the inner non-degradable portion is loaded with a biosensor. [The present invention 1018] An implantable device of the present invention 1017, wherein the biosensor is capable of measuring blood oxygen, blood glucose, tumor antigens, cholesterol, creatinine, hemoglobin A1C, insulin, glucagon, hormone levels, levels of small molecule drugs, DNA, RNA, cytokines, circulating tumor cells, among other possible diagnostic indicators. [The present invention 1019] Any of the implantable devices of claims 1001 to 1018, wherein the degradable outer housing and / or the inner non-degradable portion are patterned, e.g., the inner non-degradable portion is patterned to allow for angiogenesis once implanted. [The present invention 1020] An implantable device of the present invention 1019, wherein the patterned portion includes vascular geometries such as multiple branching blood vessels ranging in size from tens to hundreds of micrometers in diameter and having substantially smooth inner walls to minimize frictional resistance and turbulence as fluids flow through them. [The present invention 1021] The implantable device of any one of claims 1001 to 1020, wherein the device comprises multiple layers of hydrogel. [The present invention 1022] A method for monitoring a state or condition of a subject, comprising implanting a device of the present invention 1001 into said subject, said device comprising a sensor or biosensor. [The present invention 1023] A method of treating a medical condition or disease in a subject, comprising implanting a device of the present invention 1001 into said subject, wherein the device contains a therapeutic agent. [The present invention 1024] A method of producing an angioplasty device in a subject, comprising implanting a device of the present invention 1001 into said subject. [The present invention 1025] A method for creating angiogenic tissue around an implant or prosthesis in a subject, comprising implanting a device of the present invention 1001 into the subject, wherein the implant or prosthesis is a neural interface device. [The present invention 1026] A method for creating augmented tissue around an implant or prosthesis in a subject, comprising implanting a device of the present invention 1001 into the subject, wherein the implant or prosthesis is a breast implant or tissue filler. Other objects, features and advantages of the present invention will become apparent from the following detailed description, but it should be understood that the detailed description and specific examples, while indicating specific embodiments of the present invention, are given by way of illustration only, since various changes and modifications within the spirit and scope of the invention will become apparent to those skilled in the art from the detailed description. [Brief description of the drawings]

[0024] The following drawings form part of the present specification and are included to further demonstrate certain aspects of the present invention. The invention may be better understood by reference to one or more of these drawings in combination with the detailed description of specific embodiments presented herein. The patent or application file may contain at least one color drawing. Copies of this patent or patent application publication with color drawing(s) will be provided by the Office upon request and payment of the necessary fee.

[0025] [Figure 1] 1 shows a schematic diagram of an embodiment comprising a non-degradable moiety encased in a degradable polymer. [Diagram 2] Schematic diagrams of embodiments including therapeutic cells, biosensors, or other implants loaded into the non-degradable portion of the device are shown. [Diagram 3] FIG. 1 shows a schematic diagram of an embodiment including methods and materials that can be loaded into the degradable and / or non-degradable portions of the device to provide controlled release of oxygen, pro-angiogenic substances or other relevant compounds. [Figure 4] FIG. 1 shows a schematic of an embodiment including degradable moieties that can use polymer chains of various lengths to modify concentration, stiffness, and anti-fibrotic properties. [Diagram 5] 1 shows a schematic diagram of an embodiment including a vascular structure with many possible topologies. [Figure 6] 1 shows a schematic diagram of an embodiment including anchors that can hold together the non-degradable and degradable portions of the device. [Figure 7] 1A-1D show schematic diagrams of embodiments of a vascular scaffold with and without endothelial cells lining the scaffold. [Figure 8] 1 shows a hydrogel for rapid vascularization of therapeutic cell constructs. [Figure 9] We present the development of optimized shell chemistries for macrodevice development. [Figure 10] Two embodiments of the device that were tested in vivo are shown: a device with hollow vessels (left), and a device with endothelialized vessels (right). [Figure 11] FIG. 1 shows a vascular design to investigate the vascular efficiency of perfusable hydrogels. [Figure 12] A variety of device geometries for both the inner and outer housings are shown. [Figure 13] Angiogenic device for transplantation of immunoisolated islets. Pro-angiogenic cells such as endothelial cells and human mesenchymal stem cells are encapsulated in a 3D printed degradable shell, and pancreatic islets are cast into a non-degradable core of TMTD alginate. The encapsulated pro-angiogenic cells recruit vascular infiltration, and the 3D printed vascular channels help direct vascular infiltration. Over time, the vascular network remains and supplies nutrients to the immunoisolated islets. [Figure 14] Figure 14A-B. Gel preparation and design. Gels are prepared as follows: a photosensitive pre-hydrogel solution is combined with the desired cell type to encapsulate the cells in a patterned hydrogel with complex architecture and design. Endothelial cells can be seeded into the lumen of the bioprinted gel, allowing perfusion culture to enhance the angiogenic phenotype of the endothelial cells. Finally, therapeutic cells, in this case pancreatic islets, can be flowed into the chambers of the 3D printed device (Figure 14A). This results in a core-shell angiogenic device with therapeutic cells located in one chamber and pro-angiogenic cells located in the bulk gel. In this example, pancreatic islets are encapsulated in TMTD alginate (non-degradable), HUVECs and hMSCs are encapsulated in the bulk PEG / GelMA gel, and endothelial cells line the lumen (Figure 14B). [Figure 15]Biomaterial composition of angiogenic gels. Gels are prepared by stereolithography of monomers mixed with bioactive cells to create bulk encapsulation of bioactive cells in complex gel architectures. The lumen of the gel can also be seeded with endothelial cells, which can be maintained in perfusion culture to enhance the angiogenic potential of these cells. In this example, PEG diacrylate and GelMA are utilized to create gel structures to bulk encapsulate MSCs and endothelial cells to secrete angiogenic factors and aid in gel degradation. Endothelial cells are also seeded in the lumen to aid in angiogenesis. Modifying the biomaterial properties and cell composition allows for fine tuning of biomaterial properties to target desired in vivo responses. [Figure 16] Figures 16A-D. Reaction schemes for the synthesis of photoinitiators (LAP), photoreactive polymers (PEGDA and GelMA), and small molecule modified 3D printed gels. (Figure 16A) Synthesis of LAP and PEGDA. (Figure 16B) Synthesis of GelMA. (Figure 16C) Synthesis of Met-RZA15, an immunomodulatory small molecule. (Figure 16D) PEGDA and GelMA are photocrosslinked with Met-RZA15 to form a hydrogel network with pendant Met-RZA15. [Figure 17] Figure 17A-B. Hydrogel structures show an inverse correlation between collagenase degradation rate and shear stiffness. (Figure 17A) Degradation kinetics of various hydrogel formulations in 0.2 mg / mL collagenase. We observe a general trend of decreasing degradation rate as polymer concentration increases. Furthermore, the addition of PEGDA appears to have a particularly strong damping effect due to the lack of an enzymatically cleavable peptide sequence. (Figure 17B) The addition of PEGDA dramatically increases the shear stiffness when compared to pure GelMA gels. [Figure 18]Figure 18A-D. TOF-SIMS surface analysis of printed gels verifies conjugation with Met-RZA small molecule. (Figure 18A) TOF-SIMS measurement of blank gel substrate (top trace) compared to gel conjugated with Met-RZA (bottom trace). (Figure 18B) Overlay of blank gel trace from part (a) with Met-RZA gel trace in the mass range of RZA15. (Figure 18C) Overlay of blank gel trace from part (a) with Met-RZA gel trace in the mass range of Met-RZA. (Figure 18D) Quantification of TOF-SIMS data for both unmodified and small molecule modified gels. [Figure 19] Figure 19A-D. Hollow vascular channels can be printed within biocompatible 3D-printed hydrogels. (Figure 19A) A 3D model is first sliced ​​into a series of 2D patterns. (Figure 19A) The 2D patterns are sequentially projected as UV photomasks into a photosensitive polymer solution, resulting in a solid hydrogel structure with a patterned architecture. (Figure 19C) Schematic geometry of a single-channel gel, the smallest unit of a vascular network. (Figure 19D) Hollow vascular channels with increased circularity are obtained by further reducing the height of the printed layers (scale bar = 1 mm). The gel is composed of 20 wt% PEGDA (6 kDa) with 1 mg / mL FITC-dextran (150 kDa) to allow for fluorescent visualization. 3D printing of single-channel hydrogels remains a significant challenge in the field of biofabrication. Due to the nature of the overhanging circular features, the material may inadvertently clog the hollow channel in both extrusion and light-based printing methods. [Figure 20]Figure 20A-B. Fully perfusable architecture models are faithfully fabricated with projection stereolithography. We designed (Figure 20A) serpentine and (Figure 20B) branched vascular architectures using Blender (Blender Foundation, Amsterdam, Netherlands), an open-source 3D computer graphics software. We generated these engineered vascular architectures in PEGDA:GelMA (3.25 wt% 3.4 kDa PEGDA / 10 wt% GelMA) hydrogels featuring fully perfusable hollow lumens. The patency of their lumens is verified by perfusing: (Figure 20A) 1 μm red fluorescent beads (1:800 in PBS+10% glycerol) and (Figure 20B) human alveolar basal epithelial adenocarcinoma cells (A549) expressing H2B-mVenus (not shown) and mCherry. These flow throughout our engineered vascular architectures. [Figure 21]Figure 21A-F. Printed hydrogels exhibit the desired flow patterns. (Figure 21A) Fluorescent beads with a diameter of 1 μm (1:800 ratio of beads to PBS + 10% glycerol) were perfused through a hydrogel with a snake-like architecture. Video data were captured using a Nikon Eclipse Ti epifluorescence microscope (frame rate of 24.6 ms). The inlet flow rate was set at 2.0 μL / min, and the direction of flow is indicated by the white arrows. Particle image velocimetry (PIV) data on the movement of fluorescent beads were analyzed using PIVlab, an open-source application in MATLAB (Theilicke and Stamhuis, 2014). (Figure 21B) Following analysis in PIVlab, a heat map of velocity vectors was generated. The lighter areas in Figures 21A and 21D represent areas of the hydrogel containing non-moving particles, which were manually masked prior to image processing, and the arrows represent the velocity vectors. The velocity vectors are examined on a line perpendicular to the direction of flow (arrows in Figures 21B and 21E) to determine the flow characteristics within this architecture. (Figure 21C) Velocity profile data for each point along the line. Individual data points (circles) were compiled for every video frame and plotted on the same graph. The velocity was averaged across all frames for each distance along the profile line. The average velocity associated with each distance was used to generate a parabolic best-fit curve (shown in orange) (r2=0.855). (Figures 21D-F) The experiment was repeated for a hydrogel with a branched architecture. In this design, the concentration of fluorescent beads was 1:200 beads:PBS and the inlet flow rate was set to 5 μL / min. The best-fit curve for this velocity profile has an r2=0.966. [Figure 22]Figure 22A-B. 3D printed blood vessels are highly compliant under dynamic fluid cycling between systolic and diastolic pressures. (Figure 22A) Vascular expansion was measured by applying pneumatic air pressure (hydrogel composition: 20 wt% 6 kDa PEGDA, channel diameter: 1.5 mm, channel length: 20 mm). White dashed lines represent the imaged channel walls (scale bar = 1 mm). (Figure 22B) Compliance was found to be highly tunable within the polymer parameter space and directly correlated with molecular weight but inversely correlated with polymer concentration (n ≥ 4). The printed gels were connected to an open-source air pump (MillerLab Pneumatic System, 2016) to provide dynamic fluid cycling between diastolic (80 mmHg) and systolic (120 mmHg) pressures. Diameter changes were captured with an optical stereo microscope and sCMOS camera. Images were captured at rates of over 50 frames per second, allowing the maximum change in diameter to be determined with high spatial resolution. Compliance was measured as previously reported for tissue engineered blood vessels (L'Heureux et al., 2006). [Diagram 23] Figure 23A-D. A fully endothelialized vascular network covers the channel walls and maintains luminal patency. Our optimized hydrogel formulation (3.25 wt% 3.4 kDa PEGDA, 10 wt% GelMA) supports the attachment of human endothelial cells. We inject 10-30x106 cells / mL of constitutively expressing GFP human umbilical vein endothelial cells (GFP-HUVECs, Angioproteomie, Boston, MA) into the vessel inlet and allow the cells to attach to the luminal wall. Following the initial cell injection, we rotate the gel around the vessel axis by 90 degrees every 15 min for a total of 2 h. The gel is then cultured using Vasculife medium (Lifeline Technologies, LL-0003, Frederick, MD) and fluidly connects the tubing to the inlet of the vascular network via a peristaltic pump to precisely control the flow rate of 5 μl / min. Images were acquired with a Nikon Eclipse Ti epifluorescence microscope (Figures 23A, 23B, and 23D) or a Zeiss LSM 780 confocal microscope (Figure 23C). [Figure 24] More than 85% of HUVECs remain viable in the vascular network after 7 days of perfusion. The endothelialized serpentine channel was perfused at 5 μl / min for 7 days and stained for cell viability. We prepared 1 μM calcein violet AM solution (ThermoFisher Scientific) and 1 μl / mL DEAD (Ethidium Homodimer-1) component of the LIVE / DEAD imaging kit (Invitrogen, R37601) in PBS. The staining solution was injected into the channel and incubated at 37° C. for 30 minutes before imaging (Nikon Eclipse Ti epifluorescence microscope). Four regions of interest were randomly selected to manually count live and dead cells for cell viability output. [Diagram 25] Figures 25A-C. A block of TMTD alginate can be poured into a 3D printed gel. (Figure 25A) A perspective view of the printed gel. (Figure 25B) A cross-section of the printed gel, zoomed in to highlight the underlying vascular structures and holding the alginate block with a 3D printed anchor. (Figure 25C) [Figure 26] Figure 26A-D. Islets show >90% viability and stimulation index >2 after casting into vascular scaffolds. (Figure 26A) Islets in TMTD alginate were cast into 3D printed vascular scaffolds. After casting, gels were incubated with complete CMRL-1066 containing 1:200 calcein AM and 1:200 ethidium homodimer 1 for 30 minutes before being fixed and imaged. Scale bar = 100 μm. (Figure 26B) After encapsulation, islet cell viability was quantified for 5 islets. (Figure 26C) Gels were subjected to GSIS assay and showed an average stimulation index of 3.35 with n=5 independent samples. (Figure 26D) Stimulation index from GSIS assay >2. For all plots, mean + / - SD is shown. [Figure 27]Figure 27A-E. In vivo gonadal fat pad implantation demonstrates blood in the 3D printed channels and closely associated host vasculature. (Figure 27A) Surgical method of fat pad implantation of 3D printed devices. (Figure 27B) Endothelialized gel after 6 days in perfusion culture shows lining by adherent GFP-HUVECs. Scale bar = 1 mm. (Figure 27C) Gel sutured to fat pad. (Figure 27D-E) Explanted gel shows blood in the channel (white arrowhead) and closely associated host vasculature (black arrowhead). Endothelialized gel 4 with snake-shaped channels (Figure 27D) and endothelialized gel 5 (Figure 27E) before implantation (left inset). Left inset, scale bar = 1 mm. [Figure 28] Figure 28A-C. Patterned biomaterials lead to vascular invasion. 3D model of angiogenesis device for immunoisolated islets (Figure 28A, left). Bioprinted gel with perfusion-stimulated GFP endothelial cells in the S-shaped vascular channel and red endothelial cells in the majority of the gel (Figure 28A, right). We note that a monolayer of cells was obtained in the snake-like structures, a sign of endothelial maturation. Gels were sutured to the epididymal fat pad of male SCID-beige mice (Charles River Labs) (Figure 28B, left) and retrieved from the same site (Figure 28B, right). With intravenously injected fluorescent dextran, all vascular structures infiltrated into the gel were illuminated and could be visualized by confocal microscopy (Figure 28C). [Figure 29]Image analysis pipeline to identify key vascular parameters. We can leverage image analysis methods to identify key vascular parameters in hydrogels. Confocal imaging of intravenous fluorescent dextran illuminates the vascular network in the hydrogel (confocal). This image, with 16-bit color depth, is first binarized in ImageJ (Binarize). The binary image shows the location of vascular structures (black pixels) compared to the background (white pixels). This binary image is quantified to identify the percentage of blood vessels (black pixels) relative to the total number of pixels in the image. This quantifies the area fraction of the gel that is occupied by blood. The binary image is then skeletonized in ImageJ. Using the ImageJ skeletonization plugin, we can quantify the number of junctions, the number of branches, and the total length of the branches (not shown here). The junctions and branches of the sample ROI are shown below. [Diagram 30] Figure 30A-F. Designing biomaterials to achieve in vivo angiogenesis. 3D printed angiogenic devices can be fabricated with a variety of hydrogel chemistries, cells in bulk gels, and cells in channels. Three such formulations were selected and evaluated in vivo via implantation and suturing into the epididymal fat pads of SCID-beige mice (Figure 30A). Hydrogels were prepared in vitro and imaged with epifluorescence microscopy. RFP-HUVECs were present in the majority of all gels, demonstrating the presence of contracted collagen cords (GFP-HUVECs in condition 1, pre-implantation) and confluent endothelial monolayers (GFP-HUVECs in conditions 2 and 3, pre-implantation) (see Figure 30B). Prior to harvesting, CF633-conjugated dextran (Biotium) was injected via the tail vein and allowed to circulate for 15 min. Mice were euthanized and gels were fixed. Key parameters of vascular invasion were quantified (Figure 30C-F, n=4-6 replicates per group). [Diagram 31]Figure 31A-B. Kinetic study of vascular development. A key element in the vascularization of islet transplants is the kinetics of vascular development. We probed the kinetics by explanting gels at various time points. Gels were prepared in vitro and assayed for proper endothelial morphology prior to transplantation (Figure 31A, epifluorescence prior to transplantation). Gels were implanted and harvested at various time points. We note vascular invasion occurring between 4 and 8 weeks (Figure 31B). Images represent n=4-6 replicates per group. [Diagram 32]Figure 32A-G. Evaluation of fluid flow in patterned vascular networks and initial testing of single channel vascular networks in vitro. Gels with various vascular architectures can be prepared in vitro (Figure 32A). The magnitude of velocity can be quantified via tracking of fluorescent beads (Nikon Ti epifluorescence microscope, see Figure 32B). Data from the velocity profile for each point along the line. Individual data points (blue circles) were compiled for every video frame and plotted on the same graph. Velocities were averaged across all frames, for each distance, along the profile line. The average velocity associated with each distance was used to generate a parabolic best-fit curve (shown in orange) (r2=0.855, Figure 32C). To study the effects of vascular architecture on cell function and nutrient delivery, we performed in vitro assays utilizing encapsulated luciferase-expressing cells and perfusion of D-luciferin in complete medium. In the presence of ATP, oxygen, and luciferin, the luciferase enzyme is illuminated, allowing nutrient delivery and cell function to be quantified in complex hydrogels. To study nutrient delivery and cell activity, gels with various vascular structures were seeded with luciferase-expressing cells in an alginate matrix (Figure 32D). These gels were perfused with D-luciferin in complete medium (0.15 mg / mL final concentration, PerkinElmer). Data from a representative gel at 0 hours (Figure 32E) and 2 hours (Figure 32F) of perfusion by in vivo imaging (PerkinElmer IVIS KineticIII) are shown. The kinetics of cell luminescence was quantified via Living Image software and plotted in Prism 8 (Figure 32G, n=4+ / -sd). [Diagram 33]Figure 33A-D. Screening of vascular architectures for nutrient delivery. Gels with various vascular architectures were printed and the space around the vascular network was flooded with alginate containing luciferase-expressing HEK293 cells (25 million cells per mL of alginate) (Figure 33A). These hydrogels are then perfused with D-luciferin-containing medium through the vascular lumen, delivering luciferin and oxygen through the bulk gel. Through in vivo imaging (IVIS Kinetic III), we are able to quantify luminescence resulting from nutrient delivery and cellular function in the engineered hydrogels. Three different vascular architectures were prepared and analyzed for nutrient delivery (Figure 33B). Luminescence was quantified every 5 min for 120 min (Figure 33C, hexagonal and snake-shaped architectures, n=1; single channel architecture, n=9, mean + / - sem). The inventors have noted in previous studies that the hexagonal architecture had more effective nutrient delivery compared to single channel or serpentine architectures through kinetic analysis (Figure 33C) and at the study endpoint (Figure 33D). [Diagram 34]Figure 34A-D. Evaluation of high density encapsulated islets and future diabetes studies. In the diabetes study, TMTD-alginate encapsulated islets will be compared to the experimental group, i.e. hydrogel with optimized biomaterial formulation and vascular architecture for optimized nutrient delivery (Figure 34A). We also plan to perform a cell type control to control for the effect of co-implantation of islets with pro-angiogenic cells (Figure 34A). Alginate blocks were poured with 750 IEQ of fresh rat islets and transplanted into 5 STZ-induced C57Bl6 mice (Jackson Labs, Figure 34B). Mice with blood glucose levels below 200 mg / dL were considered healed. The percentage of healed mice varied between 0% and 60% over the 7-week study, indicating unstable glycemic control (Figure 34C). Upon explantation, alginate blocks were stained for islet viability (Live / Dead Assay, Thermo Fisher). Islets were found to be dead (Figure 34D). This indicates that in the absence of vasculature, densely encapsulated islets do not adequately restore normoglycemia. This study serves as a baseline for comparison with angiogenic gels. [Diagram 35] Histological examination demonstrates the formation of blood vessels in the gel. To verify that vascular invasion occurred in the bioengineered gel, samples were subjected to histological examination. 3D-printed PEG / GelMA gels, patterned with collagen cords in the lumen and containing hMSCs and HUVECs in the majority of the gel, were implanted in immunodeficient mice for 8 weeks. Gels were then harvested and subjected to histological analysis. Histological examination was performed by formalin fixation, paraffin embedding (FFPE). Gels were sectioned at 6 micron thickness and H&E stained. The entire gel was imaged (left, EVOS XL, color CCD, magnification 2X, scale = 2000 μm) as well as the inset (right, Nikon A1 RSI, water immersion 40X, Di-F3 color camera, scale = 10 μm). Red blood cells and capillary-like structures can be seen in multiple areas of the inset (right). [Diagram 36]Printing of multi-layered hydrogel constructs with varying biomaterial properties. In this example, a stereolithography 3D printing system can be utilized as follows: printing can be performed with one pre-hydrogel solution, denoted as Formulation 1, which is composed of norbornene-modified hyaluronic acid (NorHA), the thiol-terminated cell adhesion peptide CGRGDS (SEQ ID NO:2), and the dithiol-terminated degradable peptide CGPQGIWGQGCR (SEQ ID NO:1). This prints the components of the device as degradable biomaterial formulations. The print is then paused and the print solution is replaced with a solution with different properties, such as non-degradable PEG-diacrylate. This photopolymerizes in the presence of LAP to generate the second component of the device, denoted as Formulation 2. Finally, this print solution can be replaced with another solution, such as the degradable Formulation 1, to print the final component of the device. This ultimately results in a 3D printed device with various spatially separated biomaterial properties. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0026] Detailed Description of the Invention Transport and diffusion limitations remain major obstacles in the development of highly cellularized constructs. Although the challenges are well defined, the technology for developing angiogenic devices remains elusive. There is a significant medical need to develop biomaterials that overcome the challenges of eliminating FBR associated with biomaterial implantation. We have addressed this challenge through the identification of immunomodulatory small molecules by combinatorial chemistry and high-throughput screening (Vegas et al. 2016b) and verified the long-term efficacy of these small molecules in preventing fibrosis and long-term cell survival in non-human primates.

[0027] In one aspect, we have developed novel manufacturing strategies, such as biomaterial chemistry and advanced three-dimensional printing, to explicitly address the challenge of angiogenesis in engineered tissues and implantable devices (Miller, 2014). They have developed a high-throughput pipeline for the synthesis and evaluation of hundreds of materials and prototype devices, and have explored new approaches to both cell encapsulation and device fabrication. For alginate-based materials, novel lead compounds have been identified that exhibit low levels of inflammation but require further development (see below). To improve safety and efficacy, we are investigating the use of semipermeable, low-fibrotic angiogenic devices in combination with alginate sheets as containment units that facilitate complete removal of all implanted cells in the event of teratoma formation. This work lays an important foundation for functional, expandable allogeneic replacement tissues that may restore glycemic control in large animal diabetic models and ultimately in humans.

[0028] These and other aspects of the disclosure are described in detail below.

[0029] I. 3D Printable Devices A. Device Components We have previously described a 3D printing process that can fabricate 3D engineered tissues with biologically inspired design criteria, including but not limited to Murray's Law, multi-scale branching vessels with diameters of tens to hundreds of micrometers, smooth inner walls, circular cross-sections, and multiple inlets / outlets. Indeed, with optimization of print parameters, what can be fabricated is limited by what can be modeled. Furthermore, more complex designs can be fabricated, including non-uniform properties in a single layer or multiple layers. For example, grayscale photomasks with predefined gradients can be incorporated to obtain layers of varying stiffness, or layers with controlled immobilization of biomolecules and cells, while still using the same baths and solutions. Furthermore, utilizing fractal space-filling models to computationally grow vascular networks around and through existing vascular networks, or following the architecture of native tissues, can be achieved by computational growth models of even more complex and physiologically relevant 3D models. These mathematical fractal space-filling models can be derived, for example, from knot theory, Hilbert curves, and L-systems. Such mathematical fractal space-filling models for predicting idealized vascular networks include, but are not limited to, knot theory, plumber's nightmare, Peano curve, Hilbert curve, Pythagorean tree, and Brownian tree models. As an example, the plumber's nightmare model is essentially composed of two vascular ladder models interconnected by a straight vertical cylinder. Multiple plumber's nightmare models can be inserted such that they interpenetrate. The vascular ladder model is composed of one inlet and one outlet, with two horizontal cylinders connected by a diagonal cylinder, resulting in a junction between the channels.

[0030] Photopolymerizable hydrogel materials such as poly(ethylene glycol) diacrylate (PEGDA) can be crosslinked using photoinitiator systems such as lithium acylphosphinate (LAP) (Fairbanks et al. 2009), which absorbs wavelengths ranging from UV to visible light. For example, by adding low concentrations of carbon black (which can absorb light across the entire UV-visible light spectrum) or low concentrations of tartrazine (which has a peak of light absorption around 500 nm), we can limit the depth of light penetration. To quantify this process, we developed a photorheological assay to monitor the polymerization and stiffness changes of the hydrogel as a function of light dose and sample thickness. Indeed, the addition of additives such as tartrazine controls the degree of gelation of the prepolymerization mixture by affecting the kinetics of gelation and the rheological properties of the final gel. Other materials include -ene modified natural and synthetic materials that can be photopolymerized (e.g., alginate, silk, dextran, chondroitin sulfate, hyaluronic acid, cellulose, heparin, and poly(caprolactone) and multicomponent versions thereof).

[0031] To achieve complex patterning of multilayer hydrogels with high pattern fidelity, on the order of several centimeters, the light exposure during the printing process is controlled so that the light projected onto the build platform interacts primarily with the layer that undergoes either partial or complete gelation. Radical-mediated photopolymerization of hydrogels utilizes photoinitiators (molecules sensitive to a specific wavelength range that upon light absorption collapse to release free radicals that can catalyze hydrogel polymerization). This makes it essential to quantify the wavelength sensitivity of the photoinitiator. Higher concentrations of photoinitiators provide higher z-resolution by absorbing more light and limiting the penetration depth of the incident light. However, high concentrations of photoinitiators interfere with the photopolymerization reaction (the more free radicals there are, the more likely they are to annihilate each other), and high concentrations of photoinitiators are cytotoxic. Furthermore, the high xy resolution from the projector can result in light hitting previously printed layers from the z-direction, limiting the ability to form complex overhanging structures (such as those found in the vasculature) and can cause phototoxicity to trapped cells.

[0032] Thus, to achieve high-resolution printing, the present disclosure provides for the first time a photochemical means to provide high z-resolution in bioprinted tissues while maintaining high cell viability. To address the concerns outlined above, the inventors have identified a general strategy to add biocompatible materials or chemicals to the prepolymerization solution to provide higher z-resolution. The additive material is selected based on three criteria: 1) the ability to absorb light wavelengths that fully encompass the wavelength range of photosensitivity of the photoinitiator, 2) limited participation or limited inhibition of the photopolymerization reaction, and 3) biocompatibility at the desired concentration. This additive material is referred to herein as a biocompatible light-absorbing additive material suitable for controlling light transmission. Several molecules have been screened that absorb light and limit the degree of light transmission into the already formed layer. Suitable molecules absorb in the same region as the photoinitiator used in the prepolymerization solution. Examples of molecules that can control light transmission and are therefore suitable for use as biocompatible light-absorbing additive materials include carbon black, yellow food coloring, tartrazine, nanoparticles, microparticles, gold nanoparticles, riboflavin, phenol red, β-carotene, curcumin, saffron, and turmeric. Proteins can also serve as suitable biocompatible light-absorbing additive materials, so long as their peak absorption overlaps with that of the photoinitiator and matches the incident light source. Furthermore, the inventors have recognized that cells transfected or transduced with proteins that absorb in the same region as the photoinitiator, such as cyan fluorescent protein (CFP) or green fluorescent protein (GFP), can be used at high concentrations, with reduced or no additives, reducing lateral overcuring due to light-absorbing molecules present in the cells. Furthermore, the inventors' methodology allows for printing hydrogels with both horizontal and vertical channels by tightly controlling the transmission of projected light.

[0033] To address potential cell viability concerns associated with long printing times to fabricate engineered tissues with this method, cell viability can be maintained or enhanced in multiple ways. One method involves reducing the metabolic activity of cells by printing at cryogenic conditions. Similar to cryopreservation of solid organs for long-term storage of transplanted organs (typically performed at 4 °C), reducing the temperature at which cells are printed significantly reduces cell metabolism. For example, we have demonstrated that low-temperature bioprinting can be achieved utilizing a cold room (4 °C). Additionally, vitamins, growth factors, or serum can be incorporated into the prepolymerization solution to allow for an easy supply of nutrients, thus maintaining or enhancing cell viability. Importantly, photopolymerization is highly resistant to decreases in temperature, requiring only a small increase in exposure time. Additionally, reducing the temperature of the entire 3D printing device and reagents may help to reduce the heat generated from the incident light or the photopolymerization process. Additionally, molecules such as PEG and glycerol are widely used as cryoprotectants in cell culture. Thus, the use of polymers such as PEG during the cryoprinting process may improve cell viability.

[0034] Branched, multi-scale transport systems are found in all multicellular organisms. Similar to the highly complex branching structure of the vascular network, the respiratory tree is also composed of a complex branching structure to supply sufficient air to the distal regions of the lungs. It has been shown that endothelial cells can assist in the branching of the lung epithelium. However, with current fabrication techniques, structures that mimic the anatomical complexity of native lung tissue are not possible. By using 3D printing, it should be possible to generate structures that mimic the anatomical complexity of native lung tissue and vasculature. The proposed approach can make it possible to print such structures and embed endothelial and epithelial cell types in the channel lumen to mimic vascular and respiratory networks. The achievable circular cross-section allows for the development of confluent cell layers along the channel lumen. Considering the higher z-resolution under the proposed approach, the channels can more closely mimic vascular and respiratory networks. The disclosed methods and materials can allow for the fine tuning of the geometry and architecture of multiple networks. By using a fractal space-filling model similar to physiological vascular networks, this technology enables the design and fabrication of relevant 3D constructs with interpenetrating channels.

[0035] Moreover, the proposed approach can be combined with other scaffold fabrication techniques, such as porogen leaching or surface coating, to result in physiologically relevant complex constructs with modified internal microarchitecture or surface properties. Furthermore, the proposed approach can be used in the fabrication of microfluidic devices for organ-on-a-chip or human-on-a-chip applications. Furthermore, the printer can be modified to include specific sensors to ensure the printing of more precise layer thicknesses.

[0036] In certain embodiments, a prepolymerization solution is provided. The prepolymerization solution includes a photosensitive polymer having a molecular weight of more than 2,000 Daltons, a photoinitiator, and a biocompatible light-absorbing additive material suitable for controlling light transmission. In some embodiments, the prepolymerization solution can also include one or more living cells (referred to herein simply as cells). In some aspects, the prepolymerization solution can include a cryoprotectant, such as low molecular weight PEG, glycerol, ethylene glycol, sucrose, propylene glycol, trehalose, raffinose, guar gum, xanthan gum, and D-mannitol. The cryoprotectant can be permeating or non-permeating. Each of these components is described in further detail herein. The prepolymerization solution can also include a water content of 10% to about 99.5% by weight. The prepolymerization solution can include a water content of 80% to about 90% by weight. In some aspects, the prepolymerization can further include the following components: DMEM medium, serum, proteins, growth factors, thickening agents, and / or anti-aggregation agents. Such components may provide nutrients and / or neutral buoyancy to the cells in the prepolymerization solution.

[0037] In certain embodiments, photopolymers with molecular weights of more than 2,000 Daltons can be used. The photopolymers can contain at least two vinyl groups per molecule of polymer. Such vinyl groups can include acrylates, acrylamides, and methacrylates. Photopolymers that can be used include, by way of example and without limitation, poly(ethylene glycol) diacrylate (PEDGA), cell-adhesive poly(ethylene glycol), MMP-sensitive poly(ethylene glycol), poly(ethylene glycol) dimethacrylate (PEGDMA), poly(ethylene glycol) diacrylamide (PEGDAAm), gelatin methacrylate (GelMA), methacrylated hyaluronic acid (MeHA), and PEGylated fibrinogen. The photopolymers can also be modified by conjugation of cell-adhesive peptides, such as CGRGDS (SEQ ID NO:2).

[0038] In certain embodiments, photoinitiators can be used, which are molecules sensitive to a particular range of wavelengths and upon absorption of light, the molecules can break down and release free radicals to catalyze the polymerization of the hydrogel. Such photoinitiators can include, by way of example and without limitation, lithium acylphosphinate, Irgacure 2959, the Eosin Y system (Eosin Y, 1-vinyl-2-pyrrolidinone (NVP), and triethanolamine (TEA), tris(triphenephosphine)ruthenium(II), and camphorquinone, typically used with either ethyl 4-N,N-dimethylaminobenzoate or TEA, as well as isopropylthioxanthone photosensitizers. High concentrations of photoinitiators can be used to achieve improved z-resolution by limiting the penetration of the incident light, but these high concentrations can interfere with the photopolymerization reaction and be cytotoxic. Additionally, the use of high concentrations of photoinitiators limits the ability to print complex overhanging structures and can cause phototoxicity to trapped cells, as light can shine through previously printed layers.

[0039] In certain embodiments, biocompatible light-absorbing additive materials can be used. Biocompatible light-absorbing additive materials are selected based on three criteria: 1) the ability to absorb light wavelengths that completely encompass the photosensitive wavelength range of the photoinitiator, 2) limited participation or limited inhibition of the photopolymerization reaction, and 3) biocompatibility at the desired concentration. Biocompatible light-absorbing additive materials can be organic. Examples of biocompatible light-absorbing additive materials can include, but are not limited to, carbon black, yellow food coloring, tartrazine, nanoparticles, microparticles, gold nanoparticles, riboflavin, phenol red, β-carotene, curcumin, saffron, and turmeric. In addition, cells transfected or transduced with proteins that absorb in the same region as the photoinitiator, such as cyan fluorescent protein (CFP) or green fluorescent protein (GFP), can be used at high concentrations with or without reduced additives, reducing lateral overcuring due to light-absorbing molecules present in the cells. Biocompatible light-absorbing additive materials allow for the printing of hydrogels with both horizontal and vertical channels by tightly controlling the transmission of projected light, allowing for high z-resolution due to reduced light transmission into already printed layers.

[0040] In some embodiments, a hydrogel matrix is ​​provided. The hydrogel matrix can include a first tubular channel and a second tubular channel. The hydrogel matrix can be porous. The hydrogel matrix can also include a first cell type and a second cell type embedded therein. In certain aspects, the hydrogel matrix can include a first cell type embedded therein.

[0041] The hydrogel matrix may be produced in one or more layers, including, in certain embodiments, more than 1,000 layers, about 10 layers to about 2,000 layers, about 10 layers to about 1,000 layers, about 10 to about 500 layers, about 10 to about 100 layers, about 100 to about 2,000 layers, about 100 to about 1,000 layers, about 100 layers to 500 layers, about 100 layers to about 300 layers, about 500 layers to about 1,000 layers, about 500 layers to about 2,000 layers, about 1,000 layers to about 2,000 layers, and any range therebetween. In certain other embodiments, each layer may have a thickness of about 25 microns to about 100 microns, about 50 microns to about 100 microns, about 25 microns to about 50 microns, and any range therebetween. In certain other aspects, each layer may have a thickness of 50 microns. In yet other aspects, each layer may have a thickness of less than 50 microns. In some embodiments, each layer may have a thickness of 25 microns. In certain aspects, each layer may have a thickness of 100 microns. In certain aspects, the hydrogel matrix of one or more layers may include a first cell type, and the hydrogel matrix of one or more other layers may include a second cell type but not the first cell type.

[0042] One or more layers of hydrogel matrix can have cells embedded therein. In certain embodiments, one or more layers of hydrogel matrix adjacent to one or more layers of hydrogel matrix having cells embedded therein comprises extracellular matrix proteins.

[0043] The hydrogel matrix of one or more layers can be formed from a photosensitive polymer. In certain embodiments, the hydrogel matrix of one or more layers can be formed from a second photosensitive polymer. The hydrogel matrix of one or more layers can each include a first portion and a second portion. In certain embodiments, the first portion is formed from a photosensitive polymer and the second portion is formed from a second photosensitive polymer having a molecular weight of more than 2,000 Daltons. In certain embodiments, the first portion can include a first cell type embedded therein, and the second portion can include a second cell type embedded therein, the first cell type being different from the second cell type. In certain embodiments, the first portion can include a first fluorophore and the second portion can include a second fluorophore, the first fluorophore being different from the second fluorophore.

[0044] In some embodiments, the hydrogel can include a first tubular channel and a second tubular channel. In certain aspects, the first tubular channel and the second tubular channel can each include a horizontal segment that intersects with multiple layers of the bulk hydrogel matrix. The second tubular channel can interpenetrate the first channel, where interpenetration is defined as a spatial relationship between two channels, where one channel crosses a plane between two distinct portions of the other channel at least once. The tubular channels can also branch. For example, as observed in the torus knot model, the tubular channels may branch, where the tubular channels reconverge at another point within the hydrogel. However, the branched structures can also include channels that extend from the first tubular channel and / or the second tubular channel and terminate within the hydrogel. For example, dendritic structures can be designed and generated using the current approach. In certain embodiments, the tubular channel has a diameter of 300-500 microns, 500 microns or less, 400 microns or less, or 300 microns or less. The tubular channel may also be perfusable. Additionally, the tubular channel may also be expandable in response to an increase in pressure therein. The tubular channel may be lined with cells, including epithelial cells and endothelial cells. In certain aspects, the first tubular channel is lined with endothelial cells. In certain aspects, the second tubular channel is lined with epithelial cells.

[0045] The first tubular channel can also include a first tubular inlet and a first tubular outlet on the surface of the hydrogel matrix. The second tubular channel can also include a second tubular inlet and a second tubular outlet on the surface of the hydrogel matrix. The first tubular channel can include a valve or other positive feature. The tubular channel can also include a spike extending therefrom into the hydrogel matrix. The tubular channel can be filled with any suitable fluid or gas. Such fluids or gases can include, by way of example and without limitation, bodily fluids and oxygen. In certain embodiments, the first tubular channel can be filled with a fluid.

[0046] In certain other embodiments, the first tubular channel can be filled with culture medium, red blood cells, blood, urine, bile, and / or gas (such as nitrogen and / or oxygen). In certain embodiments, the second tubular channel can be filled with culture medium, red blood cells, blood, urine, bile, and / or gas (such as nitrogen and / or oxygen). The tubular channels can also be filled with one or more different fluids and / or gases.

[0047] The hydrogel of the present disclosure can include three or more tubular channels. For example, the hydrogel can include a third tubular channel and a fourth tubular channel. A tubular channel can interpenetrate a plurality of other tubular channels. For example, a third tubular channel can interpenetrate a fourth tubular channel. Similarly, a second tubular channel can interpenetrate a first tubular channel and a third tubular channel. A tubular channel network including a plurality of tubular channels can also interpenetrate at least one tubular channel network or at least one other tubular channel network. For example, a third tubular channel can interpenetrate a first tubular channel that is also interpenetrated by a second tubular channel. As another example, a third tubular channel and a fourth tubular channel can interpenetrate and interpenetrate a first tubular channel, or an interpenetrating network including a first tubular channel and a second tubular channel. In this way, complex models can be constructed that allow complex interactions between tubular channels and tubular channel networks. The above examples of multiple tubular channels are for illustrative purposes only and are not intended to limit the present disclosure.

[0048] It should be noted that the list of polymers for various features of the device may overlap, as potential modifications to any given polymer may be useful for the degradable or non-degradable portions. The degradable portions serve many purposes, including including pro-angiogenic compounds, oxygen-releasing compounds, immunomodulatory compounds, or other biologically active compounds, as well as cells (including endothelial cells). The degradable material may affect the local site where it is implanted (including, but not limited to, angiogenesis, immunomodulatory, and controlled release of other compounds) and / or impart useful benefits to the non-degradable portions of the device through these and other related means.

[0049] B. Process for Fabricating Multilayered Hydrogel Matrix Constructs In certain embodiments, a process for manufacturing a multi-layered hydrogel matrix construct is provided. First, a 3D model of the multi-layered hydrogel matrix construct is created using design software, the 3D model of the multi-layered hydrogel matrix construct includes a first computational algorithm for generating a first elongated void in the multi-layered hydrogel matrix construct to provide a first tubular channel, and a second computational algorithm for generating a second elongated void in the multi-layered hydrogel matrix construct to provide a second tubular channel, the second computational algorithm resulting in a second tubular channel interpenetrating the first tubular channel. The 3D model is then converted into a format suitable for use in 3D printing software, and a formatted 3D model is generated. The formatted 3D model is then imported into the 3D printing software, and the 3D printing software is programmed to slice the 3D model into a plurality of two-dimensional (2D) xy images. The prepolymerization solution is delivered to a vat attached to the build platform of the 3D printer, the vat is transparent, and the prepolymerization solution includes a photosensitive polymer with a molecular weight of more than 2000 Daltons and at least two vinyl groups per molecule, a light absorbing additive material to control light transmission, and a photoinitiator. The vat may also include a coating to which the hydrogel will not adhere, such as a hydrophobic coating. For example, the coating may be polydimethylsiloxane (PDMS). This allows the hydrogel to be separated from the vat without adhering. The movable Z-axis stage of the 3D printer is positioned away from the vat, the Z-axis stage includes a surface sufficient for the gelling material to adhere to it, and the distance between the surface and the inner bottom surface of the vat corresponds to the desired layer thickness of the tissue construct. A pattern is then projected onto the inner bottom surface of the vat. For example, a light source may be projected through an optical configuration such as a DLP system to generate the pattern. The layers of the multilayer hydrogel matrix construct are then polymerized.The steps of providing the prepolymerization solution, positioning the movable Z-axis stage, projecting the light source, and polymerizing the layers can be repeated one or more times, with the movable Z-axis stage moving a distance corresponding to the desired thickness of each subsequent layer, and the same or different patterns being displayed for each subsequent layer. In certain embodiments, at least the steps of providing the prepolymerization solution to the vat, positioning the movable Z-axis stage, projecting the light source through the optical arrangement, and polymerizing the layers of the multi-layered tissue construct are performed under cryogenic conditions. The cryogenic conditions can include a temperature of 4° C. The optical arrangement can include a collimator, a condenser, a filter, and a DMD array. For example, the optical arrangement can be part of a digital light processing system.

[0050] In certain embodiments, the distance is between 50 microns and 100 microns. In certain other embodiments, the distance is 50 microns. The optical configuration can include a mirror positioned to provide a pattern on the inside bottom surface of the vat. The mirror can be slid back and forth to center the projection and align the layers of the multi-layered hydrogel matrix structure.

[0051] In some embodiments, the repeating step is performed at least once to generate a second layer, and the prepolymerization solution used for the second layer includes a second photopolymer that is different from the photopolymer used to fabricate the first layer. In certain aspects, the second photopolymer can have a different molecular weight than the photopolymer. The repeating step can be performed a sufficient number of times to generate a multilayer hydrogel matrix construct having a first tubular channel and a second tubular channel.

[0052] In certain embodiments, the first computational algorithm may be derived from knot theory. In certain aspects, the first computational algorithm and / or the second computational algorithm may be a Hilbert curve. The first computational algorithm and the second computational algorithm may conform to Murray's Law. In certain aspects, the first computational algorithm may result in bifurcation of the first tubular channel. The first computational algorithm and the second computational algorithm may include a mathematical space-filling model. The mathematical space-filling model may include a plumber's nightmare, a Peano curve, a Hilbert curve, a Pythagorean tree, and a Brownian tree model.

[0053] The assembled 3D printer is an automated computer-aided 3D prototyping device that utilizes additive manufacturing to selectively pattern photosensitive biomaterials, layer by layer, to generate 3D tissue engineered constructs. The printer contains a movable Z-axis stage that is lowered onto a build platform containing a vat. The vat contains a prepolymerization solution that includes a photosensitive polymer and a photoinitiator. Attached to the Z-axis stage is the surface to which the gelling material will adhere. Additionally, the printer's base contains a 45-degree mirror that reflects a horizontal projection onto the inside bottom surface of the transparent vat. The assembled printer also contains microelectronics to automate the printing process.

[0054] To print the structure, a 3D model is created in computer-aided design (CAD) software and exported to stereolithography (.stl) format. The .stl file is imported into the software where the print parameters are entered. The software then computationally slices the 3D model into a two-dimensional (2D) xy image. This acts as a dynamic photomask and uses the entered print parameters to generate commands for controlled, automated printing. Once the transparent vat is filled with the prepolymerization mixture, the Z-axis stage is lowered into the vat so that the distance between the surface of the vat and the inside bottom surface is the desired layer thickness. The light source then projects light through an optical setup that includes a combination of collimators, condensers, filters (such as dichroic or bandpass to select specific wavelengths), and a DMD array (such as a commercial projector or pico projector) with a pattern (the first layer of the 3D model) on the inside bottom surface of the vat, resulting in a specific 3D patterned layer. After the material gels, the Z-stage automatically moves to the height of the next layer, and the process continues automatically with successive automated projection / Z-stage movements until the final 3D construct is obtained.

[0055] Additionally, the output from a projector device may not be uniform. Therefore, we homogenize the output or flat field so that the projected pattern does not have unintended non-uniform characteristics. To this end, a blank exposure was projected onto a film of phosphorescent material, and then a luminescent image was obtained. This image contains a 2D matrix of intensity values ​​that is essentially normalized and inverted in a numerical computation program such as MATLAB. The inverted image essentially acts as a filter, resulting in a more uniform output and ensuring uniform gelation of the material throughout the projected layer.

[0056] Modifications to this assembly may involve the addition of a syringe setup to automatically dispense more prepolymerization solution during the printing process, if desired. To achieve more heterogeneous builds with different materials, modifications to the assembly may include modifications to the build platform, vats, and / or Z-axis stage, allowing multiple materials to be printed automatically in this manner. For example, the build platform may be designed to accommodate multiple vats containing different materials for printing 3D hydrogels with multiple materials.

[0057] A multi-scale branching vascular network with interpenetrating airway networks can be fabricated by using a bioprinter in its basic configuration. The model is created in CAD software and exported to ".stl" format. The exported file is uploaded into the software and the print parameters are entered. Once the prepolymerization solution is prepared, it is transferred to a vat housed in the build platform. The Z-axis stage is then lowered to obtain the desired layer thickness. A series of automated projections and Z-stage translations of the 3D model result in the final 3D printed hydrogel. The 3D bioprinted hydrogel can then be removed from the Z-platform. The vascular channels of the 3D printed model can be endothelialized by perfusion of endothelial cells, and the airway channels can be epithelialized by perfusion of epithelial cells. The vascular network serves as a blood supply, delivering oxygen and nutrients to the majority of the cells in the hydrogel, while the airway network provides a means of oxygenating the vascular network.

[0058] However, when the pressure at the inlet becomes lower than that at the outlet, the valve closes, blocking the backflow of fluid through the leaflets.Negative features, such as spikes with sharp tips, can be used to control the spatial sprouting of seeded endothelial cells into the bulk hydrogel region.

[0059] Multi-layered blood vessels consisting of an outer layer of fibroblasts, an inner layer of smooth muscle cells, and an endothelialized lumen can be fabricated by using a bioprinter in a configuration that can be printed with multiple materials. Fibroblasts, smooth muscle cells, and endothelial cells are cultured separately in appropriate culture conditions. Fibroblasts and smooth muscle cells are mixed separately at the desired density with a prepolymerization solution containing a photopolymer and a photoinitiator. The prepolymerization solution containing the cells is then placed in a vat on the build platform.

[0060] Additionally, the build platform may house a vat containing acellular prepolymerization mix used to print the base of the 3D model. The Z-axis stage is lowered into the acellular prepolymerization mix at the desired layer thickness and the printing process begins. Once a sufficient base layer has been printed, the Z-axis stage retracts from the base vat and is lowered to a position of a second material containing a prepolymerization mix with smooth muscle cells at the height of the next layer. After that layer gels, the Z-axis stage moves out of the second vat, is rinsed with acellular prepolymerization mix to reduce contamination, and is lowered into a third vat containing a prepolymerization mix with fibroblasts at the same layer height as the previous projection. After this layer gels, the resulting layer consists of a pattern of two different cell types. The Z-axis stage moves out of the third vat, is rinsed, and returns to the second vat, and the process is repeated until the final multi-layered, multi-material 3D hydrogel is bioprinted. The 3D bioprinted hydrogel is then removed and endothelialized by perfusing the channels with endothelial cells.

[0061] II. Biological materials The disclosed devices can include one or more biological materials. A material is any compound, composition, conjugate, or construct that can be used to diagnose or treat a disease, disorder, condition, symptom, or the like. Examples of such materials include cells, tissues, cell products, tissue products, proteins, antibodies, vaccines, vaccine components, antigens, epitopes, drugs, salts, nutrients, buffers, acids, and bases.

[0062] The biological material for inclusion in the disclosed device can be any biological substance. For example, the biological material can be tissue, cell, small biomolecule, or biopolymer. Examples of biopolymers include nucleotides, amino acids, cofactors, and hormones. Examples of biopolymers include nucleic acids, polypeptides, proteins, and polysaccharides.

[0063] Examples of proteins include enzymes, receptors, secreted proteins, structural proteins, signaling proteins, hormones, and ligands. Any class, type, form, or particular biological material can be used in combination with other classes, types, forms, or particular biological materials.

[0064] A. Cell In some embodiments, the device may include cells or tissues (e.g., living cells or tissues), which in some embodiments are encapsulated in or coated with a polymer. In such embodiments, the surface of the polymer encapsulation or coating is modified with a moiety or compound disclosed herein. In some embodiments, the cells may include an exogenous nucleic acid encoding a therapeutic or diagnostic polypeptide.

[0065] The cell type selected for inclusion in the disclosed device depends on the desired therapeutic effect. Cells may be derived from the patient (autologous cells), another donor of the same species (allogeneic cells), or another species (xenogeneic). Xenogeneic cells are easily available, but the possibility of rejection and the risk of possible viral transmission to the patient limit their clinical application. Any of these types of cells may be derived from natural sources, stem cells, derived cells, or genetically engineered cells.

[0066] In some embodiments, the cells are genetically engineered cells that secrete a therapeutic agent, such as a protein or hormone to treat a disease or other condition. In some embodiments, the cells are genetically engineered cells that secrete a diagnostic agent. In some embodiments, the cells are stem cells, e.g., embryonic stem cells, mesenchymal stem cells, hepatic stem cells, or bone marrow stem cells.

[0067] Types of cells for inclusion in the disclosed devices include cells from natural sources (cells from heterologous tissues, cells from cadavers, and primary cells), stem cells (embryonic stem cells, mesenchymal stem cells, and induced pluripotent stem cells, derived cells (cells derived from stem cells, cells from cell lines, reprogrammed cells, reprogrammed stem cells, and cells derived from reprogrammed stem cells, etc.), and genetically engineered cells (cells genetically engineered to express a protein or nucleic acid, cells genetically engineered to produce a metabolic product, and cells genetically engineered to metabolize a toxic substance, etc.).

[0068] Types of cells for inclusion in the disclosed devices include liver cells (e.g., hepatoblasts, hepatic stellate cells, cholangiocytes, or hepatocytes), insulin producing cells (e.g., pancreatic islet cells, isolated pancreatic beta cells, or insulinoma cells), kidney cells, epidermal cells, epithelial cells, nervous system cells (neurons and glial cells (e.g., astrocytes)), ganglion cells, retinal epithelial cells, adrenal medulla cells, lung cells, cardiac myocytes, osteoblasts, osteoclasts, bone marrow cells, spleen cells, thymus cells, glandular cells, blood cells (e.g., T cells, B cells, macrophage lineage cells, lymphocytes, or monocytes), and the like. ), endocrine hormone producing cells (e.g., parathyroid cells, thyroid cells, or adrenal cells), cells of intestinal origin and other cells that act primarily to synthesize and secrete or metabolize substances, endothelial cells (e.g., capillary endothelial cells), fibroblasts (e.g., skin fibroblasts), myoblasts, keratinocytes, smooth muscle cells, progenitor cells (e.g., bone marrow progenitor cells, adipose progenitor cells, hepatic progenitor cells, endothelial progenitor cells, peripheral blood progenitor cells, or progenitor cells of muscle or skin origin), bone marrow stromal cells, cell lines (e.g., CHO cells, MDCK cells, and PC12 cells).

[0069] A particular cell type is a pancreatic islet cell or other insulin-producing cell. The hormone-producing cell can produce one or more hormones, such as insulin, parathyroid hormone, antidiuretic hormone, oxytocin, growth hormone, prolactin, thyroid stimulating hormone, adrenocorticotropic hormone, follicle stimulating hormone, luteinizing hormone, thyroxine, calcitonin, aldosterone, cortisol, epinephrine, glucagon, estrogen, progesterone, and testosterone. Genetically engineered cells are also suitable for inclusion in the disclosed device. In some embodiments, the cells are engineered to produce one or more hormones, such as insulin, parathyroid hormone, antidiuretic hormone, oxytocin, growth hormone, prolactin, thyroid stimulating hormone, adrenocorticotropic hormone, follicle stimulating hormone, luteinizing hormone, thyroxine, calcitonin, aldosterone, cortisol, epinephrine, glucagon, estrogen, progesterone, and testosterone. In some embodiments, the cells are engineered to secrete blood clotting factors (e.g., for hemophilia treatment) or to secrete growth hormone. In some embodiments, the cells are included in natural or bioengineered tissue. For example, cells for inclusion in the disclosed devices are, in some embodiments, bioartificial kidney glomeruli. In some embodiments, the cells are suitable for transplantation into the central nervous system for the treatment of neurodegenerative diseases.

[0070] In certain embodiments, cells may be included in the prepolymerization solution and / or hydrogel. Such cells may include endothelial and epithelial cell types. Such cells may be human mesenchymal stem cells (hMSCs). By way of example and not limitation, epithelial cells may include human bronchial epithelial cells (HBECs), columnar ciliated epithelial cells, mucosal cells, serous cells, basal cells, Clara cells, neuroendocrine cells, type I and type II alveolar cells, and A549 adenocarcinoma human alveolar basal epithelial cells. By way of example and not limitation, epithelial cells may be characterized based on staining for E-cadherin, surfactant proteins A, C and D, basal marker keratin 14, and TTF-1. Similarly, by way of example and not limitation, endothelial cells may include human umbilical vein endothelial cells (HUVECs), human pulmonary microvascular endothelial cells, and induced pluripotent endothelial cells. By way of example, and not limitation, endothelial cells can be characterized based on staining for platelet endothelial cell adhesion molecule (PECAM / CD31) and vascular endothelial (VE) cadherin.

[0071] The cells can be obtained directly from a donor, obtained from cell culture of cells derived from a donor, or obtained from an established cell culture line. In certain embodiments, the cells are obtained directly from a donor, washed, and combined with the polymeric material and directly implanted. The cells are cultured using techniques known to those skilled in the art of tissue culture.

[0072] Cell viability can be assessed using standard techniques such as histology and fluorescence microscopy. The function of transplanted cells can be determined using a combination of these techniques and functional assays. For example, in the case of hepatocytes, in vivo liver function tests can be performed by placing a cannula into the recipient's common bile duct. Bile can then be collected in stages. Bile pigments can be analyzed by high pressure liquid chromatography to detect underivatized tetrapyrroles, or by thin layer chromatography after conversion to azodipyrrole by reaction of diazotized azodipyrrole with ethyl anthranilate, either with or without P-glucuronidase. Diconjugated and monoconjugated bilirubin can also be measured by thin layer chromatography after alkaline methanolysis of the conjugated bile pigments. In general, the level of conjugated bilirubin increases with increasing numbers of functioning transplanted hepatocytes. Simple liver function tests (such as albumin production) can also be performed on blood samples. Similar organ function tests can be performed using techniques known to those skilled in the art, as required to determine the degree of cell function after transplantation. For example, pancreatic islet cells and other insulin producing cells can be transplanted to achieve glucose regulation through appropriate secretion of insulin. Other endocrine tissues and cells can also be transplanted.

[0073] The site or sites into which the cells are implanted are determined based on individual needs, as well as the number of cells required. For cells that replace or complement the function of an organ or gland (e.g., hepatocytes or pancreatic islet cells), the mixture can be injected into the mesentery, subcutaneous tissue, retroperitoneum, preperitoneal space, and intramuscular space.

[0074] The amount and density of cells contained in the disclosed devices will vary depending on the choice of cells and the site of implantation. In some embodiments, a single cell may be present in a concentration of 0.1×10 6 ~4×10 6 cells / ml, more specifically, 0.5×10 6 ~2×10 6The cells are present in the hydrogel capsules at a concentration of 100-10000 IE / ml. In other embodiments, the cells are present as cell aggregates. For example, islet cell aggregates (or whole islets) preferably contain about 1500-2000 cells per aggregate of 150 μm diameter, which is defined as one islet equivalent (IEQ). Thus, in some embodiments, the islet cells are present at a concentration of 100-10000 IE / ml, particularly 200-3000 IE / ml, more particularly 500-1500 IEQ / ml.

[0075] 1. Islet cells and other insulin-producing cells In certain embodiments, the disclosed compositions include islet cells or other insulin-producing cells. Methods for isolating pancreatic islet cells are known in the art. Field et al., Transplantation 61:1554 (1996); Linetsky et al., Diabetes 46:1 120 (1997). Fresh pancreatic tissue can be divided by mincing, tearing, grinding, and / or digestion with collagenase. The islets can then be isolated from contaminating cells and materials by washing, filtering, centrifugation, or selection procedures. Methods and devices for isolating and purifying islet cells are described in U.S. Patent Nos. 5,447,863, 5,322,790, 5,273,904, and 4,868,121. The isolated pancreatic cells can be optionally cultured prior to inclusion in the hydrogel capsule using any suitable method for culturing pancreatic islet cells, as known in the art. See, for example, U.S. Patent No. 5,821,121. The isolated cells can be cultured in a medium under conditions that help eliminate antigenic components. Also, the insulin-producing cells can be derived from stem cells and cell lines, and can be cells genetically engineered to produce insulin.

[0076] 2. Genetically engineered cells In some embodiments, the disclosed compositions include cells genetically engineered to produce a protein or nucleic acid (e.g., a therapeutic protein or nucleic acid). In these embodiments, the cells can be, for example, stem cells (e.g., pluripotent), progenitor cells (e.g., multipotent or oligopotent), or terminally differentiated cells (i.e., unipotent). Any of the disclosed cell types can be genetically engineered. The cells can be engineered to include a nucleic acid encoding a polynucleotide, such as, for example, an miRNA or RNAi, or a polynucleotide encoding a protein. The nucleic acid can be incorporated, for example, into the genomic DNA of the cell for stable expression, or can be incorporated, for example, into an expression vector (e.g., plasmid DNA). The polynucleotide or protein can be selected based on the disease to be treated (or effect to be achieved) and the site of transplantation or implantation. In some embodiments, the polynucleotide or protein is antineoplastic. In other embodiments, the polynucleotide or protein is a hormone, growth factor, or enzyme.

[0077] B. Non-cellular materials In some embodiments, the cells secrete therapeutically effective substances such as proteins or nucleic acids. In some embodiments, the cells produce metabolic products. In some embodiments, the cells metabolize toxic substances. In some embodiments, the cells form structural tissues such as skin, bone, cartilage, blood vessels, or muscles. In some embodiments, the cells are natural cells such as islet cells, which naturally secrete insulin, or hepatocytes, which naturally detoxify. In some embodiments, the cells are engineered to express heterologous proteins or nucleic acids and / or to overexpress endogenous proteins or nucleic acids. In some embodiments, the cells are engineered to produce new or different products, which may be the expression product of an engineered gene or another product, such as a metabolic product, produced because of an engineered gene.

[0078] Therapeutic agents that may be included in the device or engineered into the cells included in the device include, for example, thyroid stimulating hormone, beneficial lipoproteins such as Apol, prostacyclin and other vasoactive substances, antioxidants and free radical scavengers, soluble cytokine receptors (e.g., soluble transforming growth factor (TGF) receptors) or cytokine receptor antagonists (e.g., IL1ra), soluble adhesion molecules (e.g., ICAM-1), soluble receptors for viruses (e.g., CD4, CXCR4, CCR5 for HIV), cytokines, elastase inhibitors, bone morphogenetic proteins (BMPs) and BMP receptors 1 and 2, endothelial cell proliferation inhibitors (EFAs), and / or endothelial cell proliferation inhibitors (ESIs ... These include FGFs, such as FGFs, serotonin receptors, tissue inhibitors of metalloproteinases, potassium channels or potassium channel modulators, anti-inflammatory factors, angiogenic factors (vascular endothelial growth factor (VEGF), transforming growth factor (TGF), hepatic growth factor, and hypoxia inducible factor (HIF)), polypeptides with neurotrophic and / or anti-angiogenic activity (ciliary neurotrophic factor (CNTF), glial derived neurotrophic factor (GDNF), nerve growth factor (NGF), brain derived neurotrophic factor (BDNF), neurotrophin-3, nurturin, fibroblast growth factor (FGF), endostatin, ATF, fragments of thrombospondin, variants thereof, and the like. More specific polypeptides are FGFs, such as acidic FGF (aFGF), basic FGF (bFGF), FGF-1 and FGF-2, and endostatin.

[0079] In some embodiments, the active agent is a protein or peptide. Examples of protein active agents include, but are not limited to, cytokines and their receptors, and chimeric proteins comprising cytokines or their receptors (e.g., tumor necrosis factors α and β, their receptors and their derivatives), renin, lipoproteins, colchicine, prolactin, corticotropin, vasopressin, somatostatin, lypressin, pancreozymin, leuprolide, α-1-antitrypsin, clotting factors (such as factor VIIIC, factor IX, tissue factor, and von Willebrand factor), anticoagulants (such as protein C), atrial natriuretic factor; pulmonary surfactant;Plasminogen activators other than tissue-type plasminogen activator (t-PA) (e.g., urokinase), bombesin, thrombin, hematopoietic growth factors, enkephalinase, RANTES (normally expressed and secreted by T cells and regulated upon activation), human macrophage inflammatory protein (MIP-1-α), serum albumin (e.g., human serum albumin), Müllerian inhibitory factor, relaxin A chain, relaxin B chain, prorelaxin, mouse gonadotropin-related peptide, chorionic gonadotropin, microbial protein Proteins (e.g., β-lactamase), DNase, inhibin, activin, hormone or growth factor receptors, integrins, proteins A or D, rheumatoid factor, platelet-derived growth factor (PDGF), epidermal growth factor (EGF), transforming growth factors (TGFs) (e.g., TGF-a and TGF-β, including TGF-β1, TGF-2, TGF-3, TGF-4, or TGF-5), insulin-like growth factors-I and -II (IGF-I and IGF-II), des(l-3)-IGF-I (brain IGF -I), insulin-like growth factor binding proteins, CD proteins (such as CD-3, CD-4, CD-8, and CD-19), erythropoietin, bone morphogenetic factors, immunotoxins, interferons (such as interferon-α (e.g., interferon-α2A), interferon-β, interferon-γ, interferon-λ, and consensus interferon), colony stimulating factors (CSFs) (e.g., M-CSF, GM-CSF, and G-CSF), interleukins (ILs) (e.g., IL-1 through IL-10), superoxide dismutase, T cell receptors, surface membrane proteins, decay accelerating factors, transport proteins, homing receptors, addressins, fertility inhibitors (such as prostaglandins), fertility promoters, regulatory proteins, antibodies (including fragments thereof) and chimeric proteins (such as immunoadhesins), precursors, derivatives, prodrugs and analogues of these compounds, as well as pharmacologic acceptable salts of these compounds, or precursors, derivatives, prodrugs and analogues thereof. Suitable proteins or peptides may be natural or recombinant and may include, for example, fusion proteins;

[0080] In addition, examples of protein activators include CCL1, CCL2 (MCP-1), CCL3 (MIP-Iα), CCL4 (MIP-Iβ), CCL5 (RANTES), CCL6, CCL7, CCL8, CCL9 (CCL10), CCL11, CCL12, CCL13, CCL14, CCL15, CCL16, CCL17, CCL18, CCL19, CCL20, CCL21, CCL22, CCL23, CCL24, CCL25, CCL26, CCL27, CCL28, CXCL1 (KC), CXCL2 (SDFla), CXCL3, CXCL4, CXCL5, CXCL6, CXCL7, CXCL8 (IL8), CXCL9, CXCL10, CXCL11, CXCL12, CXCL13, CXCL14, CXCL15, CXCL16, CXCL17, CX3CL1, XCL1, XCL2, TNFA, TNFB (LTA), TNFC (LTB), TNFSF4, TNFSF5 (CD40LG), TNFSF6, TNFSF7, TNFSF8, TNFSF9, TNFSF10, TNFSF11, TNFSF13B, EDA, IL2, IL15, IL4, IL13, IL7, IL9, IL21, IL3, IL5, IL6, IL11, IL27, IL30, IL31, OSM, LIF, CNTF, CTF1, IL12a, IL12b, IL23, IL27, IL35, IL14, IL16, IL32, IL34, IL10, IL22, IL19, IL20, IL24, IL26, IL29, IFNL1, IFNL2, IFNL3, IL28, IFNA1, IFNA2, IFNA4, IFNA5, IFNA6, IFNA7, IFNA8, IFNA10, IFNA13, IFNA14, IFNA16, IFNA17, IFNA21, IFNB1, IFNK, IFNW1, IFNG, IL1A (IL1F1), IL1B (IL1F2), IL1Ra (IL1F3), IL1F5 (IL36RN), IL1F6 (IL36A), IL1F7 (IL37), IL1F8 (IL36B), IL1F9 (IL36G), IL1F10 (IL38), IL33 (IL1F11), IL18 (IL1G), IL17, KITLG, IL25 (IL17E), CSF1 (M-CSF), CSF2 (GM-CSF), CSF3 (G-CSF), SPP1, TGFB1, TGFB2, TGFB3, CCL3L1, CCL3L2, CCL3L3, CCL4L1,CCL4L2, IL17B, IL17C, IL17D, IL17F, AIMP1(SCYE1), MIF, Areg, BC096441, Bmp1, Bmp1O, Bmp15, Bmp2, Bmp3, Bmp4, Bmp5, Bmp6, Bmp7, Bmp8a, Bmp8b, Clqt nf4, Ccl21a, Ccl27a, Cd70, Cerl, Cklf, Clcf1, Cmtm2a, Cmtm2b, Cmtm3, Cmtm4, Cmtm5, Cmtm6, Cmtm7, Cmtm8, Crlf1, Ctf2, Ebi3, Edn1, Fam3b, Fas1, Fgf2 , Flt31, GdflO, Gdfll, Gdfl5, Gdf2, Gdf3, Gdf5, Gdf6, Gdf7, Gdf9, Gml2597, Gml3271, Gml3275, Gml3276, Gml3280, Gml3283, Gm2564, Gpi1, Grem1, Grem 2, Grn, Hmgb1, Ifnal1, Ifnal2, Ima9, Ifnab, Ifne, 1117a, 1123a, 1125, 1131, Iltifb, Inhba, Lefty1, Lefty2, Mstn, Nampt, Ndp, Nodal, Pf4, Pglyl, Prl 7dl, Scg2, Scgb3al, Slurp1, Spp1, Thpo, TnfsflO, Tnfsfl1, Tnfsfl2, Tnfsfl3, Tnfsfl3b, Tnfsfl4, Tnfsfl5, Tnfsfl8, Tnfsf4, Tnfsf8, Tnfsf9, Tslp, Vegfa, Wnt1, Wnt2, Wnt5a, Wnt7a, Xcl1, epinephrine, melatonin, triiodothyronine, thyroxine, prostaglandins, leukotrienes, prostacyclin, thromboxane, islet amyloid polypeptide, Müllerian inhibitory factor or ho luminol, adiponectin, corticotropin, angiotensin, vasopressin, arginine vasopressin, atriopeptin, brain natriuretic peptide, calcitonin, cholecystokinin, cortistatin, enkephalin, endothelin, erythropoietin, follicle-stimulating hormone, galanin, gastric inhibitory polypeptide, gastrin, ghrelin, glucagon, glucagon-like peptide-1, gonadotropin-releasing hormone, growth hormone-releasing hormone, hepcidin, human chorionic gonadotropin, human placental lactogen, growth hormone, inhibin, insulin,Somatomedins, leptin, lipotropin, luteinizing hormone, melanocyte-stimulating hormone, motilin, orexin, oxytocin, pancreatic polypeptide, parathyroid hormone, pituitary adenylate cyclase-activating peptide, prolactin, prolactin-releasing hormone, relaxin, renin, secretin, somatostatin, thrombopoietin, thyrotropin, thyrotropin-releasing hormone, vasoactive intestinal peptide, androgens, androgens, acid maltase (α-glucosidase), glycogen phosphorylase, glycogen debranching enzyme, phosphofructokinase, phosphoglycerate kinase, phosphoglycerate mutase, lactate dehydrogenase, carnitine palmitoyltransferase, carnitine, and myoadenylate deaminase.

[0081] The hormones contained in the disclosed devices or produced from the cells contained in the devices can be any hormone of interest. Examples of endocrine hormones include antidiuretic hormone (ADH) (produced by the posterior pituitary gland, targets the kidneys, and affects water balance and blood pressure), oxytocin (produced by the posterior pituitary gland, targets the uterus, breasts, and stimulates uterine contractions and milk production), growth hormone (GH) (produced by the anterior pituitary gland, targets the body's cells, bones, and muscles, and affects growth and development), prolactin (produced by the anterior pituitary gland, targets the breasts, and maintains milk production), growth hormone releasing hormone (GH) (produced by the anterior pituitary gland, targets the breasts, and maintains milk production), and growth hormone releasing hormone (GH). Thyroid-stimulating hormone (RH) (GH-releasing hormone, produced in the arcuate nucleus of the hypothalamus), thyroid-stimulating hormone (TSH) (produced by the anterior pituitary gland, targets the thyroid gland and regulates thyroid hormones), thyrotropin-releasing hormone (TRH) (produced by the hypothalamus, stimulates the release of TSH and prolactin from the anterior pituitary gland), adrenocorticotropic hormone (ACTH) (produced by the anterior pituitary gland, targets the adrenal cortex and regulates adrenal cortical hormones), follicle-stimulating hormone (FSH) (produced by the anterior pituitary gland, targets the adrenal cortex and regulates adrenal cortical hormones), , targets the ovaries / testes and stimulates egg and sperm production), luteinizing hormone (LH) (produced by the anterior pituitary gland, targets the ovaries / testes and stimulates ovulation and release of sex hormones), luteinizing hormone-releasing hormone (LHRH) (also known as gonadotropin-releasing hormone (GnRH), a trophic peptide hormone synthesized in and released by GnRH neurons in the hypothalamus and involved in the release of FSH and LH), thyroxine (produced by the thyroid gland, targets somatic cells and regulates metabolism), calcitonin (produced by the thyroid gland, targets somatic cells and regulates metabolism), parathyroid hormone (produced by the adrenal cortex, targets the adrenal cortex, and lowers blood calcium), parathyroid hormone (produced by the parathyroid gland, targets bone matrix, and raises blood calcium), aldosterone (produced by the adrenal cortex, targets the kidneys, and regulates fluid balance), cortisol (produced by the adrenal cortex, targets body cells, and dampens the immune system and stress response), epinephrine (produced by the adrenal medulla, targets the heart, lungs, liver, and body cells, and primarily affects the "fight or flight" response),Glucagon (produced by the pancreas, targets the liver, body, and increases blood glucose levels), insulin (produced by the pancreas, targets somatic cells, and decreases blood glucose levels), estrogen (produced by the ovaries, targets the reproductive system, and affects puberty, menstruation, and gonadal development), progesterone (produced by the ovaries, targets the reproductive system, and affects puberty, menstrual cycle, and gonadal development), testosterone (produced by the adrenal glands, testes, targets the reproductive system, and affects puberty, gonadal development, and sperm).

[0082] In some embodiments, the protein is a growth hormone (such as human growth hormone (hGH), recombinant human growth hormone (rhGH), bovine growth hormone, methionine-human growth hormone, des-phenylalanine human growth hormone, and porcine growth hormone), insulin, insulin A chain, insulin B chain, and proinsulin, or a growth factor (vascular endothelial growth factor (VEGF), nerve growth factor (NGF), platelet-derived growth factor (PDGF), fibroblast growth factor (FGF), epidermal growth factor (EGF), transforming growth factor (TGF), and insulin-like growth factor-I and insulin-like growth factor-II (IGF-I and IGF-II)).

[0083] III. Alginate Immunoisolation A fundamental barrier to successful cell therapy is the lack of biocompatible implantable devices, including in the context of implantable devices. Implanted biomaterials trigger the foreign body reaction (FBR), a cascade of inflammatory events and wound healing processes that leads to fibrosis (partition formation) and subsequent implant failure (12-14). This reaction has been described for many materials, from naturally occurring polymers to synthetic materials (15-17). Thus, a significant medical need exists to develop biomaterials that overcome this key challenge of eliminating the FBR.

[0084] To address this issue, we contemplate using immunoisolation alginate to create a protective coating for the cells contained in the device of the present application. We used combinatorial chemistry and high-throughput screening to identify immunomodulatory small molecules (Vegas et al., 2016b) and validated the long-term efficacy of these small molecules in preventing fibrosis and long-term cell survival in non-human primates.

[0085] The present disclosure provides, in certain embodiments, a device comprising a degradable shell and a non-degradable core. In certain embodiments, the non-degradable core can contain encapsulated therapeutic cells and / or biosensors, and the degradable shell serves as a scaffold for vascular growth and enhances blood flow to the cells and / or biosensors. The inventors have applied modern 3D bioprinting technology to induce vascularization of immunoisolated therapeutic cell constructs for long-term survival and functionality, allowing for higher density therapeutic cell engraftment.

[0086] One embodiment provides the use of modified alginate polymer in the construction of a cell-containing device. The modified alginate polymer can be a polymer of any desired molecular weight. The weight average molecular weight of the alginate is preferably 1,000 to 1,000,000 Daltons, more preferably 10,000 to 500,000 Daltons, as determined by gel permeation chromatography.

[0087] The modified alginate polymer can include any ratio of mannuronic acid monomers, guluronic acid monomers, and covalently modified monomers. In some embodiments, 2.5%, 5%, 7.5%, 10%, 12%, 14%, 15%, 16%, 18%, 20%, 22%, 24%, 25%, 26%, 28%, 30%, 32.5%, 35%, 37.5%, 40%, 45%, 50%, 55%, or 60% of the monomers in the modified alginate polymer are covalently modified monomers. More than 10%, more than 20%, or more than 30% of the monomers in the modified alginate polymer are covalently modified monomers.

[0088] Modified alginate polymers can be produced incorporating covalently modified monomers with a variety of different hydrogen bonding potentials, hydrophobicity / hydrophilicity, and charge states. The inclusion of covalently modified monomers in an alginate polymer alters the physiochemical properties of the alginate polymer. Thus, the physiochemical properties of alginate can be tailored to the desired application by selectively incorporating covalently modified monomers.

[0089] For example, the glass transition temperature (T g ) may be altered by the incorporation of covalently modified monomers. In some embodiments, the modified alginate polymer powder has a T of greater than 50° C., 60° C., 65° C., 70° C., 75° C., 80° C., 85° C., 90° C., 95° C., 100° C., 105° C., 110° C., 115° C., 120° C., 125° C., 130° C., 135° C., 140° C., 145° C., 150° C., 160° C., 175° C., 190° C., or 200° C. as measured by differential scanning calorimetry (DSC). g has.

[0090] The hydrophobicity / hydrophilicity of alginates can be altered by incorporating hydrophobic and / or hydrophilic covalently modified monomers. In certain embodiments, the modified alginate polymer comprises one or more hydrophobic covalently modified monomers. The relative hydrophobicity / hydrophilicity of modified alginates can be quantitatively assessed by measuring the contact angle of a water droplet on a film of modified alginate polymer using a goniometer. In some embodiments, the modified alginate has a contact angle of less than 90 degrees (i.e., it is hydrophilic). In certain embodiments, the modified alginate has a contact angle of more than 90 degrees (i.e., it is hydrophobic). In some embodiments, the modified alginate has a contact angle of more than 95 degrees, 100 degrees, 105 degrees, 110 degrees, 115 degrees, or 120 degrees.

[0091] In embodiments used for cell encapsulation, the modified alginate polymer contains Ca 2+ , Sr 2+ , or Ba 2+ and the like to form hydrogels.

[0092] In some embodiments, the modified alginate polymer forms a hydrogel such that the fluorescence intensity measured using the high throughput hydrogel formation assay described herein is greater than 10,000, 15,000, 20,000, 25,000, 30,000, 35,000, 40,000, 45,000, 50,000, or 55,000. In certain embodiments, the modified alginate polymer forms a hydrogel such that the fluorescence intensity measured using the high throughput hydrogel formation assay described herein is greater than 15,000. In certain embodiments, the modified alginate polymer forms a hydrogel such that the fluorescence intensity measured using the high throughput hydrogel formation assay described herein is greater than 15,000. In certain embodiments, the modified alginate polymer forms a hydrogel such that the fluorescence intensity measured using the high throughput hydrogel formation assay described herein is greater than 15,000, 55,000, preferably 20,000 to 55,000, more preferably 25,000 to 55,000.

[0093] In an embodiment used for cell encapsulation, the modified alginate polymer forms a modified hydrogel that is porous enough to prevent immune cell ingress into the gel matrix while at the same time allowing nutrients, waste products, and hormones and / or proteins secreted by the encapsulated cells to freely diffuse into and out of the capsule. The porosity and surface area of ​​the modified alginate hydrogel can be measured using BET analysis. Prior to BET analysis, solvents and volatile impurities are removed by prolonged heating of the modified alginate gel under vacuum. The hydrogel sample is then cooled under vacuum, for example by liquid nitrogen, and analyzed by measuring the volume of gas (typically N2, Kr, CO2, or Ar gas) absorbed into the hydrogel at a particular pressure. Analysis of the physical adsorption of gases at different pressures is used to characterize the total surface area and porosity of the gel formed by the modified alginate polymer. A particular method for determining the porosity of a hydrogel is BET analysis.

[0094] In certain embodiments, the modified alginates form modified hydrogels that are porous enough to prevent immune cell infiltration into the gel matrix while at the same time allowing nutrients, waste products, and hormones and / or proteins secreted by the encapsulated cells to freely diffuse into and out of the capsule. In some embodiments, the porosity of the hydrogels formed by the modified alginate polymers is increased by 5%, 10%, 15%, or 20% compared to the porosity of the hydrogels formed from unmodified alginate polymers. In alternative embodiments, the porosity of the hydrogels formed by the modified alginate polymers is decreased by 5%, 10%, 15%, or 20% compared to the porosity of the hydrogels formed from unmodified alginate polymers.

[0095] In certain embodiments used for cell encapsulation, the modified alginates are biocompatible. The biocompatibility of the modified alginates can be quantitatively determined using a fluorescence-based in vivo biocompatibility assay described in Example 5. In this assay, cathepsin activity was measured using an in vivo fluorescence assay to quantify the foreign body response to the modified alginates.

[0096] In some embodiments, the modified alginate polymer is biocompatible such that the normalized fluorescence response relative to unmodified alginate measured using the in vivo biocompatibility assays described herein is less than 100%, 95%, 90%, 85%, 80%, 75%, 70%, 65%, 60%, 55%, 50%, 45%, or 40%. In certain embodiments, the modified alginate polymer induces a lower foreign body response than unmodified alginate. This is shown by a normalized fluorescence response relative to unmodified alginate of less than 100%. In some embodiments, the modified alginate polymer is biocompatible such that the normalized fluorescence response relative to unmodified alginate measured using the in vivo biocompatibility assays described herein is less than 75%, more preferably less than 65%, and most preferably less than 50%.

[0097] The modified alginate can be chemically modified to any desired density of modification as described herein. The density of modification is the average number of modifications (i.e., attached compounds) per given weight, volume, or surface area of ​​a capsule or product containing the modified alginate. In general, a density equal to or greater than a threshold density can provide beneficial effects, such as reduced foreign body reaction. In some embodiments, high density is not required. Without being bound to any particular theory of operation, it is believed that the chemical modification signals, indicates, or is identified by one or more immune system or other body components, resulting in beneficial effects, such as lower foreign body reaction. In some embodiments, a lower density of modification can be effective for this purpose.

[0098] Useful densities include at least, less than, about, or greater than 1, 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, 20, 25, 30, 35, 40, 45, 50, 55, 60, 65, 70, 75, 80, 85, 90, 95, 100, 110, 120, 130, 140, 150, 160, 170, 180, 190, 200, 210, 220, 230, 240, 250, 260, 270, 280, 290, 300, 310, 320, 330, 340, 350, 360, 370, 380, 390, 400, 410, 420, 430, 440, 450, 460, 470, 480, 490, 500, 510, 520, 530, 540, 550, 560, 570, 580, 590, 610, 620, 630, 640, 650, 660, 670, 680, 690, 700, 710, 720, 730, 740, 750, 760, 770, 780, 790, 800, 850, 900, 9 130, 140, 150, 160, 170, 180, 190, 200, 210, 220, 230, 240, 250, 260, 270, 280, 290, 300, 320, 340, 360, 380, 400, 420, 440, 460, 480, 500, 550, 600, 650, 700, 750, 800, 850, 900, and 1000 modifications. All ranges defined by any pair of these densities are also specifically contemplated and disclosed.

[0099] In some embodiments, the density of modifications on a surface, or portion of a surface, of the capsule or product that will come into contact with fluids, cells, tissues, other components, or combinations thereof, of the subject's body when the product is administered to a subject (e.g., implanted within the subject's body) is greater than the density of modifications on other surfaces of the product.

[0100] Density can also be expressed in terms of the concentration of surface modifications as measured by X-ray photoelectron spectroscopy (XPS), a surface-sensitive quantitative spectroscopic technique that measures the elemental composition to the parts per thousand of elements present within a material.

[0101] XPS spectra are obtained by irradiating a material with an X-ray beam and simultaneously measuring the kinetic energy and number of electrons emitted from the top 0-10 nm of the material being analyzed. By measuring all elements present at the surface, the percentage of elements that originate from the surface modification can be calculated. This can be accomplished, for example, by taking the percentage of nitrogen (and / or other elements in the surface modification) in the total elemental signal measured. Many substrates and materials from which capsules or products are formed contain very little nitrogen, so nitrogen is a useful indicator of surface modification. For convenience, the percentage of an element used to indicate surface modification can be expressed as the percentage of surface modification. Also, for convenience, the percent surface modification is sometimes referred to as the concentration of surface modification.

[0102] Useful percent surface modification include concentrations of about 0.1, 0.2, 0.3, 0.4, 0.5, 0.6, 0.7, 0.8, 0.9, 1, 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, 20, 25, 30, 35, 40, 45, 50, 55, 60, 65, 70, 75, 80, 85, 90, 95, and 100 percent surface modification. All ranges defined by any pair of these concentrations are also specifically contemplated and disclosed.

[0103] Useful percent surface modification also include concentrations of less than 0.1, 0.2, 0.3, 0.4, 0.5, 0.6, 0.7, 0.8, 0.9, 1, 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, 20, 25, 30, 35, 40, 45, 50, 55, 60, 65, 70, 75, 80, 85, 90, 95, and 100 percent surface modification. All ranges defined by any pair of these concentrations are also specifically contemplated and disclosed.

[0104] Useful percent surface modification also include concentrations of 0.1, 0.2, 0.3, 0.4, 0.5, 0.6, 0.7, 0.8, 0.9, 1, 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, 20, 25, 30, 35, 40, 45, 50, 55, 60, 65, 70, 75, 80, 85, 90, 95, and 100 percent surface modification. All ranges defined by any pair of these concentrations are also specifically contemplated and disclosed.

[0105] In one embodiment, alginates can be used to form capsules. The capsules are particles with an average diameter of about 150 μm to about 5 cm. The disclosed capsules can be formed from crosslinked hydrogels. Other than the encapsulated material, the capsules can be formed, for example, of crosslinked hydrogels alone, can have a crosslinked hydrogel core surrounded by one or more polymer shells, can have one or more crosslinked hydrogel layers, can have a crosslinked hydrogel coating, or can have a combination thereof. The capsules can have any shape suitable for cell encapsulation, for example. The capsules can contain one or more cells dispersed in a crosslinked hydrogel, thereby "encapsulating" the cells. Certain capsules are formed from or contain one or more of the disclosed modified alginates.

[0106] The average diameter of the capsules is about 150 μm to about 8 mm. The capsules can have any average diameter from about 150 μm to about 5 cm. In particular, the capsules have an average diameter of more than 1 mm, more particularly 1.5 mm or more. In some embodiments, the capsules can be as large as about 8 mm in diameter. For example, the capsules can be in the size range of about 1 mm to 8 mm, 1 mm to 6 mm, 1 mm to 5 mm, 1 mm to 4 mm, 1 mm to 3 mm, 1 mm to 2 mm, 1 mm to 1.5 mm, 1.5 mm to 8 mm, 1.5 mm to 6 mm, 1.5 mm to 5 mm, 1.5 mm to 4 mm, 1.5 mm to 3 mm, or 1.5 mm to 2 mm.

[0107] The rate at which molecules necessary for cell survival enter the capsule and the rate at which therapeutic agents and waste products exit the capsule membrane can be selected by adjusting the permeability of the capsule. The permeability of the capsule can also be altered to limit the entry of immune cells, antibodies, and cytokines into the capsule. In general, as shown by the examples, known methods of forming hydrogel capsules can produce capsules whose permeability limits the entry of immune cells, antibodies, and cytokines into the capsule. Different cell types have different metabolic requirements, so the permeability of the membrane can be optimized based on the cell type encapsulated in the hydrogel. The diameter of the capsule is an important factor that affects both the immune response to the cell capsule as well as the mass transport across the capsule membrane.

[0108] Increasing awareness of the parameters that drive fibrosis in vivo has been applied to the analysis of the performance of modified alginates. Intraperitoneal (IP) implantation of modified alginate capsules revealed that modified alginates, when crosslinked using conditions defined for unmodified alginates, can result in abnormally shaped capsules. These abnormally shaped capsules can complicate the implementation and interpretation of IP implanted modified alginate capsules. To improve capsule morphology, a formulation method was developed for use with modified alginate microparticles blended with small amounts of high molecular weight alginate. Particles prepared from this mixture resulted in particles with improved morphology and stability.

[0109] Unmodified alginates usually have a weight average molecular weight of about 50,000 Daltons to about 500,000 Daltons. However, unmodified alginates having a molecular weight can also be used. In some embodiments, the weight average molecular weight is about 50,000 to about 250,000 Daltons, more preferably about 50,000 to about 150,000 Daltons. In some embodiments, the weight average molecular weight is about 100,000 Daltons.

[0110] In other embodiments, one or more additional hydrogel-forming polymers are used in combination with or instead of unmodified alginate.Such polymers are known in the art.Examples include, but are not limited to, PEG, chitosan, dextran, hyaluronic acid, silk, fibrin, poly(vinyl alcohol) and poly(hydroxyethyl methacrylate).

[0111] In some embodiments, the alginate is composed of β-D-mannuronic acid (M) and aL-guluronic acid (G) linked together. In some embodiments, the alginate is a high guluronic acid (G) alginate. In some embodiments, the alginate is a high mannuronic acid (M) alginate. In some embodiments, the ratio of M:G is about 1. In some embodiments, the ratio of M:G is less than 1. In some embodiments, the ratio of M:G is greater than 1.

[0112] Particles prepared from a mixture of modified and unmodified alginates produced a more homogenous population of microparticles in terms of shape and size as assessed by scanning electron microscopy (SEM).

[0113] In some embodiments, the hydrogel capsule can have any suitable shape. Useful shapes include sphere, spherical shape, spheroid, spheroid shape, ellipsoid, ellipsoid shape, stadiumoid, stadium shape, disk, disk shape, cylinder, cylinder shape, rod, rod shape, cube, cube shape, cuboid, cuboid shape, torus, torus shape, and flat and curved surfaces. The coated or coated products, devices, and surfaces can have any of these shapes, or any shape suitable for the product or device.

[0114] A sphere, an ellipsoid of revolution, and an ellipsoid are shapes having surfaces that can be defined by rotation about each of three mutually perpendicular axes a, b, and c, which are circles, ellipses, or a combination thereof. In the case of a sphere, the three axes are of the same length. In the case of an oblate spheroid (also called a rotationally flattened ellipsoid), the axis lengths are a = b > c. In the case of a prolate spheroid (also called a rotationally prolate ellipsoid), the axis lengths are a = b < c. In the case of a triaxial ellipsoid (also called a scalene ellipsoid), the axis lengths are a > b > c. A stadium shape is a rotationally stadium-shaped form. A cylinder is a rotationally rectangular form that rotates about its major axis. A disk is a cylinder that has a diameter larger than its height and is flattened. A rod is an elongated cylinder whose major axis is more than 10 times its diameter.

[0115] The terms "spherical shape", "ellipsoidal shape of revolution", "ellipsoidal shape", "stadium-shaped shape", "cylindrical shape", "rod-shaped shape", "cubic shape", "rectangular parallelepiped shape", and "toroidal shape" each refer to an object having a surface that roughly forms a sphere, an ellipsoid of revolution, an ellipse, a stadium shape, a cylinder, a rod, a cube, a rectangular parallelepiped, or a torus. Beyond the perfect or classical form of the shape, the spherical shape, the ellipsoidal shape, the ellipsoidal shape, the stadium-shaped shape, the cylindrical shape, the rod-shaped shape, the cubic shape, the rectangular parallelepiped shape, and the toroidal shape can have waves and undulations.

[0116] Generally, a spherical shape is an ellipsoid (of the averaged surface) whose semi-axes and major axes are within 10% of each other. The diameter of a sphere or spherical shape is the average diameter, e.g., the average of the semi-axes and major axes. Generally, a spheroidal shape is an ellipsoid (of the averaged surface) whose semi-axes and major axes are within 100% of each other. The diameter of a spheroid or spheroidal shape is the average diameter, e.g., the average of the semi-axes and major axes. Generally, an ellipsoidal shape is an ellipsoid (of the averaged surface) whose semi-axes and major axes are within 100% of each other. The diameter of an ellipsoid or ellipsoidal shape is the average diameter, e.g., the average of the semi-axes and major axes. Generally, a stadium-shaped shape is a stadium (of the averaged surface) whose opposite semi-axes and major axes are within 20% of each other. The diameter of a stadium-shaped or stadium-shaped shape is the average diameter, e.g., the average of the semi-axes and major axes. Alternatively, the size of a stadium-shaped or stadium-shaped shape can be given as the average of the major axes. Generally, a cylindrical feature is a cylinder (of the averaged surface) with the semi-axes and major axes being within 20% of each other. The diameter of a cylinder or cylindrical feature is the average diameter, i.e. the average of the semi-axes and major axes.

[0117] Alternatively, the size of a cylinder or cylinder-like shape can be given as the average of the major axes. Generally, a rod-like shape is a rod (of the averaged surface) whose semi-axes and major axes are within 10% of each other. The diameter of a rod or rod-like shape is the average diameter, such as the average of the semi-axes and major axes. Alternatively, the size of a rod or rod-like shape can be given as the average of the major axes. Generally, a cubic shape is a cube (of the averaged surface) whose sides are within 10% of each other. The diameter of a cube or cubic shape is the average side length. Generally, a rectangular shape is a rectangular prism (of the averaged surface) whose corresponding sides are within 10% of each other. The diameter of a rectangular prism or rectangular prism shape is the average side length.

[0118] In general, a torus-like shape is a torus (of the averaged surface) with semi-axes and major axes within 10% of each other. The diameter of a torus or torus-like shape is the average diameter, i.e. the average of the semi-axes and major axes. Alternatively, the size of a torus or torus-like shape can be given as the diameter of the entire ring.

[0119] A "flat surface" refers to a continuous area of ​​a surface that has zero curvature that is greater than 5%. An "acute angle" refers to a location on a surface where the tangent to the surface changes by more than 10% over a distance that is less than or equal to 2% of the circumference of the surface. All edges, vertices, grooves, and edges on a surface form acute angles.

[0120] Certain capsules can be made of a biocompatible material, have a diameter of at least 1 mm and less than 10 mm, have a spheroidal shape, and have one or more of the following additional characteristics: pores on the surface of the capsule greater than 0 nm and less than 10 μm, the surface of the capsule is neutral or hydrophilic, the surface curvature at all points on the surface of the capsule is at least 0.2 and not greater than 2, and the surface of the capsule is free of flat faces, sharp angles, grooves, or edges. In general, the capsules induce less of a fibrotic response after implantation than the same capsule lacking one or more of these features present on the capsule.

[0121] In some embodiments, the capsules are provided as a preparation, and at least 30%, 35%, 40%, 45%, 50%, 55%, 60%, 65%, 70%, 75%, 80%, 85%, 90%, 95% or 100% of the capsules in the preparation have a shape characteristic described herein, e.g., have a spheroidal shape or have a surface curvature of at least 0.2 to 2.0 at all points on their surface.

[0122] In some embodiments, the hydrogel capsules have an average diameter of more than 1 mm, particularly 1.5 mm or more. In some embodiments, the capsules can be as large as 8 mm in diameter. For example, the hydrogel capsules can be 1 mm to 8 mm, 1 mm to 6 mm, 1 mm to 5 mm, 1 mm to 4 mm, 1 mm to 3 mm, 1 mm to 2 mm, 1 mm to 1.5 mm, 1.5 mm to 8 mm, 1.5 mm to 6 mm, 1.5 mm to 5 mm, 1.5 mm to 4 mm, 1.5 mm to 3 mm, 1.5 mm to 2 mm, 2 mm to 8 mm, 2 mm to 7 mm, 2 mm to 6 mm, 2 mm to 5 mm, 2 mm to 4 mm, 2 mm to 3 mm, 2.5 mm to 8 mm, 2.5 mm to 7 mm, 2.5 mm to 6 mm, 2.5 mm to 5 mm, 2.5 mm to 4 mm, 2.5 mm to 3 mm, 3 mm to 8 mm, 3 mm to 7 mm, 3 mm to 6mm, 3mm-5mm, 3mm-4mm, 3.5mm-8mm, 3.5mm-7mm, 3.5mm-6mm, 3.5mm-5mm, 3.5mm-4mm, 4mm-8mm, 4mm-7mm, 4mm-6mm, 4mm-5mm, 4.5mm-8mm, 4.5mm-7mm, 4.5mm-6mm, 4.5mm-5mm, 5mm-8mm, 5mm-7mm, 5mm-6mm, 5.5mm-8mm, 5.5mm-7mm, 5.5mm-6mm, 6mm-8mm, 6mm-7mm, 6.5mm-8mm, 6.5mm-7mm, 7mm-8mm, or 7.5mm-8mm in size range. In some embodiments, the capsule has an average diameter or size of 1mm-8mm. In some embodiments, the capsule has an average diameter or size of 1mm-4mm. In some embodiments, the capsules have an average diameter or size of 1 mm to 2 mm. In some embodiments, the capsules are provided as a preparation, wherein at least 30%, 35%, 40%, 45%, 50%, 55%, 60%, 65%, 70%, 75%, 80%, 85%, 90%, 95%, or 100% of the hydrogel capsules in the preparation have a diameter in a size range described herein.

[0123] Suitable alginates and modified alginates are described in US2016 / 0030360A1, US2017 / 0239397A1, WO2012 / 167223, and WO2017 / 075631, each of which is incorporated herein by reference.

[0124] IV. Therapeutic uses Therapeutic applications include the treatment of cancer, such as when immunomodulatory proteins can be produced by therapeutic cells encapsulated in the microdevice. Therapeutic production of blood factors such as FVIII or FIX can be used to alleviate hemophilia. These proteins can be produced by therapeutic cells encapsulated in the microdevice.

[0125] Certain embodiments of the present disclosure may include the treatment of diabetes, a global epidemic that afflicts over 285 million people worldwide (Veiseh et al., 2015). The current standard of care for people suffering from type I diabetes (TID) consists of daily administration of exogenous insulin and a strict regimen of diabetic diet along with blood glucose monitoring, but these individuals still face the challenge of nasty side effects of the disease that are in part the result of daily compliance issues (Veiseh et al., 2015; Pickup, 2012). Furthermore, the process by which beta cells in the pancreatic islets of Langerhans secrete insulin in response to fluctuations in blood glucose is imperfectly simulated by regular insulin injections (Veiseh et al., 2015; Robertson, 2004). Thus, the "holy grail" of diabetes treatment research is attempting to achieve insulin independence in TID patients by transplantation of exogenous donor islet tissue or cells. A corollary of this goal is to develop methods to eliminate the need for lifelong immunosuppression, which is required as a result of allo- or xenotransplantation and has its own complications, in order to fully implement this approach in clinical practice ( Hirshberg, 2007 , Pagliuca et al., 2014 ).

[0126] Recently, human pluripotent stem cells (hPSCs) were reported to differentiate into functional pancreatic β-cells in vitro, providing for the first time a route to an unlimited supply of human insulin-producing tissue (Pagliuca et al., 2014; Vegas et al., 2016a). Methods to mitigate the need for lifelong immunosuppression are essential to enable widespread clinical implementation of this new tissue source (Hirshberg, 2007; Shapiro, 2012; Vogel, 2014). Cell or tissue encapsulation could theoretically overcome the need for immunosuppression by protecting the graft from host immune recognition and subsequent graft rejection (Dolgin, 2014; Jacobs-Tulleneers-Thevissen, 2013).

[0127] V. Biosensors Another aspect of the present disclosure relates to biosensors. Both eukaryotic and prokaryotic cells have biological systems for binding and recognizing many analytes. Membrane proteins relay information through cells, and engineered cellular biosensors can provide feedback or analytical responses that can be detected via medical imaging and / or in blood.

[0128] Various noncellular biosensors, including continuous glucose monitors, rely on consistent access to the bloodstream. Vascularization of the biosensor can provide higher temporal resolution and higher signal levels for the analyte. Long-term acceptance of the biosensor via anti-fibrotic materials can provide long-term efficacy.

[0129] VI. Exemplary Device Architectures, Components, and Variations Thereof Embodiments of the present invention include advanced 3D bioprinting-based modular devices for angiogenesis and for immunoisolation of therapeutic cells, biosensors, and other implants that may benefit from angiogenesis. Reference is now made to FIG. 1. An exemplary embodiment shows a degradable polymer encasing a non-degradable polymer. We developed a simple planar vascular structure within a 3D printed hydrogel, which has an empty cavity for seeding therapeutic cells encapsulated in an immunoisolated alginate hydrogel. This approach allows for independent control of vascular architecture, immunoisolation, and immunomodulation (prevention of fibrotic responses). Our 3D printed chambers are biocompatible and can be tailored for biodegradation through individual optimization of the hydrogel formulation (a combination of poly(ethylene glycol) diacrylate and MMP-degradable peptide-methacrylate). The biodegradable chambers are a provisional matrix that promotes patterned anastomosis and host incorporation while at the same time supporting immunoisolated therapeutic cells within the non-degradable alginate during the early stages of implantation and remodeling in vivo. This expandable design can be layered or engineered into multi-layered structures of blood vessels, ultimately enabling the restoration of pancreatic function, the provision of clotting factors or enzymes to patients suffering from genetic diseases, and the effective vascularization of biosensors.

[0130] A. Advanced 3D Bioprinting of Biocompatible Hydrogels Previous studies in bioengineering have sought to provide mass transfer to engineered tissues by incorporating various cell types (Finkenzeller et al., 2006; Lovett et al., 2009; Wenger et al., 2005), growth factors, or microfluidic networks (Andrae et al., 2008; Bauer et al., 2005), but most of these techniques lack the ability to generate vascularized tissues with complex hierarchical vasculature while providing quantitative non-invasive metrics across the cellular environment. Three-dimensional (3D) printing holds great promise in the field of tissue engineering as this technique allows researchers to control all (x, y, z) coordinates within the printed volume.

[0131] Here, we introduce a new approach in the fabrication of biocompatible hydrogels, namely the automated fabrication of angiogenic hydrogels using a low-cost stereolithography system developed by us. We demonstrate for the first time that highly defined angiogenic structures (x, y, z feature resolution of 50 μm) can be generated with vascular channels typically on the order of 0.5-1 mm in diameter, and that dozens of scaffolds can be readily generated per day, for the studies described in this proposal. Here, we seek to engineer living tissues and maintain control over the location of therapeutic cells and the composition of the surrounding extracellular matrix (ECM). For the studies described herein, we utilize photopolymerizable synthetic hydrogels with a water content of >80%, based on poly(ethylene glycol) diacrylate (PEGDA) and MMP-sensitive peptides. PEGDA is a hydrophilic and biocompatible material, facilitating functionalization with acrylate-TMTD small molecules to modulate the host immune response (39, 40).

[0132] B. Advanced vascular topology to support therapeutic cells The perfusable vascular architecture stabilizes the homeostasis and function of therapeutic cells in the 3D gel, and its convection allows for the culture of high densities of therapeutic cells, maximizing the therapeutic potential of such engineered tissues. The inventors' hydrogel material was synthesized in-house and fabricated using custom 3D bioprinting with a stereolithography system, enabling the inventors to generate diverse 3D vascular networks within the hydrogel ex vivo.

[0133] To improve convection and diffusion to the trapped therapeutic cells, we further develop a hydrogel chamber precisely positioned next to the vascular structure and containing the trapped therapeutic cells. Our 3D printing stereolithography system allows the construction of an empty cavity for the seeding of a second alginate hydrogel with trapped therapeutic cells, and importantly, the fabrication of the underlying perfusable vascular structure. We hypothesize that by facilitating the delivery of nutrients and oxygen to the therapeutic cells, their function in vivo will be improved.

[0134] Hydrogel chambers can also be fabricated using a PEGDA:peptide MA mixture, i.e., an optimal formulation of 10 wt% peptide MA and 3 wt% PEGDA (3.4 kDa), that supports the attachment of human endothelial cells to the vessel wall while maintaining vascular patency (Figure 8E). Endothelial cells are injected into the vascular network at 25e6 cells / mL in media, and the gel is rotated 180 degrees around the vessel axis every 15 min for 4 h to allow endothelial cells to attach to all vessel surfaces. Gels are cultured overnight up to 7 days to form a stable endothelial monolayer.

[0135] Endothelial monolayers can be characterized for healthy cobblestone morphology using confocal microscopy with staining for VE-cadherin, ZO-1, and CD31. Endothelial barrier function can be assessed using histamine and thrombin as previously described (Chrobak et al., 2006). We have previously shown that patterned endothelial networks can induce anastomoses with host vasculature to form hybrid human / mouse vessels. We therefore further characterize the degradation kinetics of chambers made from a mixture of PEGDA:peptide MA (containing 0.4 mg / mL collagenase). We hypothesize that vascular development will be promoted around primary therapeutic cells encapsulated in alginate.

[0136] C. Incorporation of therapeutic cells encapsulated within TMTD alginate gel into angiogenic devices A typical cell encapsulation procedure involves the following steps: Clusters of therapeutic cells are cultured in Ca-free Krebs-Henseleit buffer (4.7 mM KCl, 25 mM HEPES, 1.2 mM KH2PO4, 1.2 mM MgSO4). -4 The cells are then washed with 100x 7x HO, 135 mM NaCl, ~pH 7.4, ~290 mOsm). After washing, the therapeutic cells are centrifuged again and all supernatant is aspirated. The therapeutic cell pellet is then resuspended in TMTD alginate solution at various cluster densities and synthesized as previously described (Vegas et al., 2016a). The solution is poured into the chamber of the degradable 3D printed hydrogel described above and BaCl -2 Crosslink the gel using gelling solution. 2+ , Sr 2+ Any divalent cation, including aqueous solutions, can be used for crosslinking. Immediately after crosslinking, the device is incubated in cell culture medium until just before implantation.

[0137] D. In Vivo Performance We have previously shown that patterned endothelial cells can anastomose with host vasculature and form hybrid human / mouse vessels after 14 days (Baranski et al., 2013). Based on our previous work, we expect device anastomosis to occur within 2 weeks in endothelialized devices, and after 4 weeks, we expect the PEGDA:peptide MA gel network to be completely degraded.

[0138] VII.Alternative uses A. Use of the Device for Therapeutic Cells and / or Biosensors Biosensors, including continuous glucose monitors, as well as therapeutic cells, would benefit from a strong vascular supply. Now, referring to FIG. 2, therapeutic cells, biosensors, or other grafts can be loaded into the non-degradable portion of the device. This device can be used to improve vascularization of any kind of graft, resulting in better data acquisition and longer-term performance of biosensors, as well as superior viability and functionality of therapeutic cells.

[0139] B. Sustained release of compounds Sustained release of oxygen may improve long-term therapeutic cell viability. Referring now to FIG. 3, methods and materials that provide controlled release of oxygen and / or pro-angiogenic compounds can be loaded into the degradable and / or non-degradable portions of the device. Oxygen-releasing compounds such as CaO2 and pro-angiogenic compounds such as VEGF can promote vascularization of the large and complex scaffolds of the degradable portions of the device. CaO -2 is an oxygen generating compound. Sustained release of pro-angiogenic compounds may improve vascularization of the device. Crystallized VEGF and cells that constitutively secrete VEGF are possible ways to provide sustained release of these compounds. Methods and materials for sustained release of compounds can be incorporated into the degradable and / or non-degradable parts of the device.

[0140] C. Modification of Degradation Profile and Stiffness of Degradable Moieties The degradation profile and stiffness of the gel can be controlled by the following parameters: PEGDA of different molecular weights (e.g., 6 kDa) can also be used to modify the degradation profile. See now Figure 4. The degradable moieties can be modified using polymer chains of various lengths. This allows for a variety of biodegradation profiles. Larger or more complex structures may require fine-tuned degradation profiles for successful in vivo effects. Vary the amount of PEGDA, GelMA, photoinitiator, and photoabsorber (Figure 9). Devices with very large and complex vascular scaffolds may require slower degradation. The strength and duration of UV crosslinking can affect the degradation and stiffness of the gel.

[0141] D. Anchors of various shapes for connecting the degradable and non-degradable portions of the device As structures become more complex, the anchors provide a connection between the degradable and non-degradable portions. This allows for surgical manipulation and implantation without structural failure of the device. Now, referring to FIG. 6, the anchors can hold the non-degradable and degradable portions of the device together. Complex structures and large non-degradable portions may require a physical bond to maintain their integrity for surgical manipulation and implantation. The anchors successfully connect both segments.

[0142] E. The stackable nature of the device The stackable nature of this device allows for layered vasculature. Larger therapeutic constructs can be developed by simply stacking multiple devices. Now, refer to Figure 18A-F. 3D printed PEGDA hydrogel chambers are shown for rapid vascularization of therapeutic cell constructs. Figure 8A shows that the printed chambers house primary islets (green) suspended in alginate (blue) and have an underlying patterned vascular network (red). The chambers can be stacked to allow for scale-up. Figure 8B shows the 3D printed chambers, demonstrating perfusable vasculature (red) and hydrogel anchors that physically capture the alginate hydrogel containing the islets. Scale bar = 1 mm. (Unpublished results). Figure 8C shows a close-up of the highlighted area from Figure 8B. Showing the hydrogel anchors and underlying perfusable vasculature (red). Scale bar = 1 mm. (Unpublished results). Figure 8D shows a 250 μm cross-section, illustrating the alginate seeding volume (blue) and the underlying vascular structures (red). Scale bar = 1 mm. (Unpublished results). Figure 8E shows that endothelial cells injected into the vasculature rapidly line the vessels as seen in the top view (left) and form hollow lumens (confocal virtual section, right) (Unpublished results). In Figure 8F, we have previously shown that after 14 days, patterned endothelial cells anastomose with the host vasculature and hybrid human / mouse blood vessels can be formed (Baranski et al., 2013).

[0143] F. Modification of Immunomodulatory Properties and Inducible Degradation of Degradable Moieties As shown in Figure 9, the hydrogel precursors can be varied. The amount and / or type of antifibrotic small molecules can be altered for different levels of immunomodulatory activity. Cell adhesion peptides can be modified to improve endothelial cell adhesion to the hydrogel vascular scaffold. The peptides of the "MMP-sensitive PEG-diacrylate" can be selected to allow for other inducible degradation mechanisms (oxide-driven degradation, pH-driven degradation, ultrasound-driven degradation, thermally-driven degradation, etc.).

[0144] G. Advanced 3D Bioprinting of Biocompatible Hydrogels Here, we introduce a novel approach in the fabrication of biocompatible hydrogels, namely the automated fabrication of angiogenic hydrogels using a low-cost stereolithography system developed by us. We demonstrate for the first time that highly defined angiogenic structures (x,y,z feature resolution of 50 μm) can be generated with vascular channels typically on the order of 0.5-1 mm in diameter for the studies described in this proposal, and that dozens of scaffolds can be readily generated per day. Here, we seek to engineer living tissues and maintain control over the location of beta cells and the composition of the surrounding extracellular matrix (ECM). We utilize photopolymerizable synthetic hydrogels with a water content of over 80%, based on poly(ethylene glycol) diacrylate (PEGDA) and MMP-sensitive peptides, for the studies described herein. PEGDA is a hydrophilic, biocompatible, FDA-approved material that has found broad utility in exploring fundamental cell-matrix interactions and exhibits ease of functionalization with acrylate-TMTD small molecules to modulate the host immune response ( Burdick and Anseth, 2002 , Lutolf et al., 2003 ).

[0145] H. Advanced vascular topology supporting 3D β cell culture To improve convection and diffusion to the trapped beta cells, we further develop hydrogel chambers precisely positioned next to the vascular structures and containing trapped beta cells. Our 3D printing stereolithography system allows the construction of an empty cavity for the seeding of a second alginate hydrogel with trapped primary islets, and importantly, the fabrication of the underlying perfusable vascular structures. We hypothesize that by facilitating the delivery of nutrients and oxygen to the islets, their function in vivo will be improved. Hydrogel chambers can also be fabricated using a PEGDA:peptide MA mixture, i.e., an optimal formulation of 10 wt% peptide MA / 3 wt% PEGDA (3.4 kDa), that supports the attachment of human endothelial cells to the vessel wall while maintaining vessel patency (Figure 8E). Endothelial cells are injected into the vascular network at 25e6 cells / mL in medium, and the gel is rotated around the vessel axis by 180 degrees every 15 min for 4 h to allow the attachment of endothelial cells to all vessel surfaces. Gels are cultured overnight up to 7 days to form a stable endothelial monolayer. We characterize the endothelial monolayer for healthy cobblestone morphology using confocal microscopy with staining for VE-cadherin, ZO-1, and CD31. Endothelial barrier function is assessed using histamine and thrombin as previously described (Chrobak et al., 2006). We have previously shown that patterned endothelial networks can induce anastomoses with host vasculature to form hybrid human / mouse vessels. We hypothesize that patterned vasculature will support host engraftment. Therefore, we further characterize the degradation kinetics of chambers made from a mixture of PEGDA:peptide MA (containing 0.4 mg / mL collagenase). We hypothesize that vascular development will be promoted around primary islets encapsulated in alginate.

[0146] I. Methods for incorporating islets encapsulated within TMTD alginate gels into angioplasty devices A typical cell encapsulation procedure involves the following steps: Syngeneic mouse islet clusters purchased from CellTrans Inc. Chicago, IL, were centrifuged at 1,400 rpm for 1 min and encapsulated in Ca-free Krebs-Henseleit buffer (4.7 mM KCl, 25 mM HEPES, 1.2 mM KH2PO4, 1.2 mM MgSO4). 2- The islets are then washed with 500 mM NaCl, 0.5×7H2O, 135 mM NaCl, approximately pH 7.4, approximately 290 mOsm). After washing, the islet cells are centrifuged again and all the supernatant is aspirated. The islet pellet is then resuspended in TMTD alginate solution at various cluster densities, 500 clusters per 0.5 mL of alginate solution, as previously described (Vegas et al., 2016a). The solution is poured into molds and the gel is crosslinked using BaCl2 gelling solution. Immediately after crosslinking, the encapsulated islet clusters are washed four times with 50 mL of CMRL1066 modified medium and cultured overnight in spinner flasks at 37 °C, immediately prior to transplantation. Since some islet clusters are inevitably lost during the encapsulation process, the total number of encapsulated clusters is recounted after encapsulation.

[0147] After encapsulation, cell viability and functionality are characterized. Cell viability is measured using a fluorescence microscopy-based assay, where cells are stained with a combination of fluorescein diacetate (live) and propidium iodide (dead). The percentage of cell viability is determined by sampling 100 islet clusters and analyzing the ratio of live and dead cells. We further perform an in vitro glucose-stimulated insulin secretion assay (Vegas et al., 2016a). EXAMPLES

[0148] VIII. Examples The following examples are included to demonstrate preferred embodiments of the present disclosure. It should be understood by those skilled in the art that the techniques disclosed in the following examples represent techniques that the inventors have found to work well in the implementation of the present disclosure, and therefore can be considered to constitute preferred modes for its implementation. However, those skilled in the art should understand in light of the present disclosure that many changes can be made in the specific embodiments disclosed and still obtain similar or similar results without departing from the spirit and scope of the present disclosure.

[0149] Hypothetical Example 1 Assessment of in vivo biocompatibility and vascular function of the angiogenesis device. We have previously shown that patterned endothelial cells can anastomose with host vasculature and form hybrid human / mouse vessels after 14 days (Baranski et al., 2013). We use healthy non-diabetic C57 / BL6 nu / nu murine T cell deficient mice to assess the biocompatibility and angiogenesis of the hydrogel chambers described above. We also assess the biocompatibility of the gels with immunomodulatory chemicals. For these studies, we use non-diabetic C57 / BL6 nude mice (Jackson Laboratories, Bar Harbor Maine). Using three replicates of the device (TMTD alginate only, TMTD alginate in chamber gel with hollow vessels, and TMTD alginate chamber gel with endothelialized vessels), we test in vivo for 2 and 4 weeks.

[0150] We validate the in vivo biocompatibility and vascular function assessment of hydrogel projection 3D printed hydrogel chambers with perfusable vascular networks to support high density 3D beta cell culture. Three replicates of the device are tested in vivo using healthy C57 / BL6 nu / nu mice, TMTD alginate only (N=5 mice per time point) (left), TMTD alginate in chamber gels with hollow vessels (N=5 mice per time point) (middle), and TMTD alginate in chamber gels with endothelialized vessels (N=5 mice per time point) (right). Two recovery time points, 2 weeks and 4 weeks, are used for evaluation.

[0151] Healthy C57 / BL6 nu / nu mice are used to test empty (islet-free) angiogenic devices in vivo. Device replicates outlined in Figure 8A-F are implanted into sites in the peritoneal cavity of C57 / BL6 nu / nu mice (Vegas et al., 2016a). In this study, TMTD alginate alone is incorporated as a control non-angiogenic system, and TMTD alginate in a chamber gel with hollow vessels is a second control system in which anastomosis with host endothelial cells is not promoted. Each mouse receives a single device graft. N=5 mice per device replicate and per harvest time point. Mice are sacrificed at 2 and 4 weeks post-implantation. Based on previous studies, we expect device anastomosis should occur within 2 weeks with endothelialized devices, and after 4 weeks, we expect the PEGDA:peptide MA gel network to have been completely degraded. At these defined 2- and 4-week time points, mice are sacrificed and devices are carefully explanted by microdissection to avoid disrupting the vascular network around the remaining alginate gel block. Recovered tissues and gels are preserved in paraformaldehyde and further processed using standard histological processing techniques.

[0152] Ex vivo analysis of biocompatibility and vascular network integrity around empty (island-free) angiogenic devices via histology. The following evaluations are then performed on the retrieved devices to assess angiogenesis and biocompatibility: 1) quantification of fibrotic overgrowth (FO) by phase contrast and confocal fluorescence microscopy, 2) FACS analysis to assess the amount of inflammatory cells (CD68-stained macrophages, Ly6Gr+-stained neutrophils, smooth muscle actin-stained myofibroblasts, and CD31 endothelial cells) in single cell suspensions derived from: i) IP washes of the retrieved solution, ii) cells dissociated from the surface of the retrieved device, and 3) histological analysis of the device to assess FO thickness, tissue inflammatory state, and angiogenesis. Tissue samples are examined for evidence of fibrosis and rejection. Masson's Trichrome staining is used to monitor collagen deposition associated with fibrosis. Additionally, sections are immunostained for fibroblasts, neutrophils, and macrophages, as well as endothelial cells, using methods previously developed by the inventors (Vegas et al., 2016a; 2016b; Veiseh et al., 2015a).

[0153] The density of vascularization around the device was verified by histology, and we expect that more than 90% of the surface area of ​​the printed channels should be colonized by host endothelial cells. Anastomosis of the endothelial cells with the host vasculature took 1-2 weeks, and the matrix PEGDA:PeptideMA needed to remain intact while anastomosis occurred. Based on previous experience, the PEGDA:PeptideMA ratio should last for 2 weeks. However, if the gel degrades too slowly, we can simply change the ratio of these two monomers to achieve the desired degradation rate. Increasing the PEGDA:PeptideMA ratio will decrease the degradation rate, while decreasing the ratio will increase the degradation rate. For example, chambers made only from PEGDA are non-degradable in vivo.

[0154] In vivo study of islet cell-loaded vascular networks in streptozotocin-induced (STZ-induced). Using a syngeneic transplantation model, we evaluate the efficacy of islet-loaded TMTD alginate gel on angiogenic devices in STZ-treated C57 / BL6 nu / nu mice. Angiogenesis around the TMTD gel should promote improved cell viability and enhanced glucose-stimulated insulin secretion kinetics.

[0155] The vascular scaffold is seeded with endothelial cells. The vascular scaffold can be seeded with endothelial cells that secrete VEGF to improve the in vivo response of the device. Now, refer to FIG. 7. Endothelial cells can be used to line the vascular scaffold. The endothelial cells secrete VEGF to promote angiogenesis of the device. Different sources of endothelial cells (off-the-shelf and patient-derived) or no cells at all can be used. This may result in different engraftment of the device in vivo. The device can be seeded with endothelial cells of any origin, so maximum safety can be ensured by using patient-derived endothelial cells.

[0156] Reference is now made to FIG. 10. Two iterations of the device can be tested in vivo: a device with hollow vessels (left), and a device with endothelialized vessels (right). The iterations can be used to validate the in vivo function of islet-loaded TMTD alginate chambered with patterned vascular networks in diabetic nude mice. Two iterations of the device are tested in vivo using STZ-treated C57 / BL6 nu / nu mice, syngeneic islet-loaded TMTD alginate in chamber gel with hollow vessels (N=5 mice) (left), and syngeneic islet-loaded TMTD alginate in gel chamber with endothelialized vessels (N=5 mice at each time point) (right).

[0157] We encapsulate syngeneic islet cells (500 per device) in TMTD alginate gels of angiogenic devices to evaluate cellular morphology, viability, and glucose-responsive insulin secretion. The functionality and viability of islets encapsulated within the gel assemblies are monitored using a combination of live-dead cell viability assays and tests of insulin production in response to glucose challenge.

[0158] Transplanted syngeneic islets are loaded into the vascular network of STZ-B6 nu / nu mice and blood glucose levels are monitored daily for up to 100 days. Six diabetic nude mice are IP implanted with syngeneic islet-loaded TMTD alginate in chamber gels with hollow blood vessels, and six diabetic nude mice are IP implanted with syngeneic islet-loaded TMTD alginate in chamber gels with endothelialized blood vessels. Mice are monitored for 100 days. Blood glucose levels are measured daily. At 14, 28, and 100 days after implantation, oral glucose tolerance tests (OGTTs) are performed on animals to monitor the in vivo kinetics of both blood glucose correction and c-peptide secretion. Healing is determined as a blood glucose level below 220 mg / dL for the entire monitoring period of 1 week and 100 days after implantation. Efficacy of encapsulation is then determined based on long-term survival and kinetics of glucose-responsive insulin secretion. Once the in vivo euglycemic experiment is finished, transplantation is performed and histopathological analysis is performed. Potential causes of graft failure will be determined, and islets will be examined for viability (live / dead assay), functionality (ex vivo glucose loading assay), and biocompatibility (immunostaining to characterize potential immune responses and fibrosis).

[0159] The efficacy of TMTD encapsulated islets in angiogenic devices is determined in a diabetic nude mouse model. Previous studies have shown that 500 islets per mouse are sufficient to achieve euglycemic correction in STZ STZ-B6 nu / nu mice. However, if glucose correction is not achieved in diabetic mice after one week, we will implant devices loaded with larger amounts of islets (1000 and 2000) into additional mice.

[0160] We develop and validate next-generation 3D culture of beta cells within 3D gels containing perfusable vasculature. We hypothesize that the perfusable vascular architecture will stabilize beta cell homeostasis and function in the 3D gel, and that convection will allow for the culture of high densities of therapeutic cells, maximizing the therapeutic potential of such engineered tissues. The hydrogel material was synthesized in-house and fabricated using custom 3D bioprinting with a stereolithography system, thus enabling the inventors to generate diverse 3D vascular networks within ex vivo hydrogels. As shown in Figure 11A-F, we screen diverse vascular topographies to identify lead formulations that support the highest density of cell viability and functionality in vivo.

[0161] Variations in Vascular Topology. Vascular topology can have a profound effect on the cultivation of high density therapeutic cells. Now, referring to FIG. 5, 3D vascular structures can be printed in many possible topologies. The topology of the 3D vascular structure can have a significant impact on the long-term viability and performance of the encapsulated therapeutic cells and biosensors. The 3D printing approach disclosed herein allows for the rapid synthesis of highly diverse and complex vascular structures. As shown in FIG. 11A-F, the 3D printing approach allows for the development of highly complex vascular topologies.

[0162] Modifying the antifibrotic properties and stiffness of alginate. The alginate used in the non-degradable portion can be modified in various ways, and a selected antifibrotic small molecule can be conjugated to the alginate at a selected dose. Furthermore, the stiffness of the alginate can be modified by changing the ratio of modified alginate to SLG100 or by changing the cross-linking buffer (Ca 2+ or Sr 2+ This may improve the biomechanical response at different implantation sites.

[0163] Figures 11A-F show vascular designs to investigate vascular efficiency in perfusable hydrogels. 3D vascular geometries are computationally generated using a space-filling algorithm with iterative complexity. Figure 11A shows a sample vascular topology growth algorithm in 2D format for clarity. Figure 11B shows that we also identified computational space-filling algorithms as good candidates for building perfusable vascular topologies. Figure 11C shows that 3D vascular-inspired architectures that follow Murray's law for scaling of daughter branch diameters can be converted to 3D-printed topologies by Boolean subtraction from bounding volumes. Figures 11D-F show vascular topologies fabricated in perfusable hydrogels shown by macrophotography or μCT analysis, as labeled. We further demonstrate that 3D vascular networks can enable interpenetrating networks, as shown in Figures 11D-E (red, blue).

[0164] Materials and Methods for Examples 2-3 GelMA, PEGDA, and LAP synthesis. LAP was prepared as previously described (Fairbanks et al., 2009) using equimolar amounts of dimethylphenylphosphinite and 2,3,6-trimethylbenzoyl chloride overnight at room temperature under argon. In the next step, poly(ethylene glycol) diacrylates (PEGDA) of 3.4 kDa, 6 kDa, 10 kDa, 20 kDa, and 35 kDa were synthesized by reaction between poly(ethylene glycol) acryloyl chloride in the presence of triethylamine overnight (Figure 16A) (Miller et al., 2010). Gelatin methacrylate (GelMA) was synthesized by adding methacrylic anhydride (0.1 mL per 1.0 g gelatin) dropwise to a mixture of gelatin (type A, bloom 300) dissolved in 100 mM carbonate / bicarbonate buffer (pH = 9.5) at 50 °C for 4 h ( Shirahama et al., 2016 ).

[0165] Fabrication of hydrogels by projection stereolithography. The integrated hydrogels were fabricated with projection stereolithography (pSLA), which creates 3D volumes by sequentially crosslinking photosensitive polymers layer by layer via the projection of a photomask (Arcaute, K. Ann. Biomed. Eng., 34:1429-41, 2006). The inventors' lab developed a custom pSLA 3D printer, consisting of a commercial digital light processing (DLP) projector, stepper motors and axes for the z-stage, a RepRap Arduino Megaboard (RAMBo), and 3D printed housing components fabricated from consumer grade poly(lactic acid) (PLA) plastic filament using an Ultimaker 2 (Ultimaker, Netherlands). The complete setup has been described previously (Grigoryan, B. Science 364(6439):458-464, 2019). Hydrogels are fabricated layer-by-layer by bonding to a glass slide surface on a z-stage that is immersed in a poly(dimethylsiloxane) (PDMS)-coated (6:1 base:catalyst) vat of prepolymer solution. The XY build size is 64x40mm and the XYZ resolution is 50µm, determined by the projector optics (XY) and the light absorber used (Z). The light source is 405nm and is interfaced to a computer to control the motors, heating elements and photomask projection.

[0166] A pre-hydrogel mixture was prepared containing 3.25 wt% 3.4 kDa PEGDA, 10 wt% GelMA, 17 mM LAP, 10% glycerol, and 2.255 mM tartrazine. Tartrazine (yellow food coloring FD&C Yellow 5, E102) was used as a light absorber, which we have previously established acts as a low-toxicity dye and is easily washed off, resulting in a transparent hydrogel for ease of imaging (Grigoryan, B. Science 2019). When cells are encapsulated in the bulk of the hydrogel, the desired concentration and co-culture ratio of cells are included and resuspended in the pre-hydrogel solution (5x106 Endothelial cells ~1x10 6 Mesenchymal stem cells). The pre-hydrogel mixture is loaded into the PDMS coated dish and the build platform is lowered to the position of the first layer. Creation Workshop software (World Wide Web at envisionlabs.net / ) can control the printer with GCode commands sent for vertical movement of the z-stage, projection of the photomask sequence, and heating control of the heating elements. As GelMA is sensitive to thermal gelation, the vat is kept at 37 °C to hold the aqueous pre-hydrogel mixture until crosslinked. The height of each layer is set to 50 μm and the exposure time is 19.5 mW / cm. 2 The output of 100 µL was adjusted for each batch of GelMA (typically 10-15 s per layer). After the printed construct was completed, the 3D hydrogel was detached from the glass slide by applying a razor between the hydrogel and the slide and allowed to swell overnight in PBS with multiple washes.

[0167] Cell culture and endothelialization for long-term perfusion. Human umbilical vein endothelial cells (HUVECs) were cultured in VascuLife medium and supplements (LifeLine Technologies, Frederick, MD) at 37 °C and 5% CO2. All HUVECs used for the endothelialized channels were passaged 4–8 times. The hydrogel was placed within a custom PLA gasket to serve as the perfusion chamber, fitting snugly against the gel, as well as placing 18-gauge needle tips at the inlet and outlet. Prior to placing the hydrogel, the chamber was attached to a glass slide, sterilized with ethanol for 15 min, and allowed to dry completely. Once the hydrogel was placed within the chamber and 18-gauge needles were attached at the inlet and outlet, the top of the chamber was sealed with another glass slide. This setup allowed for fluidically sealed perfusion through the lumenized architecture.

[0168] HUVECs were loaded into the lumen of the hydrogel microchannel at a concentration of 30e6 cells / mL using a Normject 1mL syringe connected to an 18-gauge flexible needle at the inlet of the channel. Similarly, an 18-gauge flexible needle was catheterized at the outlet to fluidically seal the channel. After injecting the suspension, the cells were allowed to attach to the bottom of the cylindrical surface for 15 minutes. The gel was then rotated 90 degrees every 15 minutes for the next 6 hours to establish a cylindrical coating of attached HUVECs. During the perfusion setup, unattached cells were washed away with fresh medium.

[0169] A high-precision multichannel peristaltic pump ISMATEC (Cole Parmer) was used to maintain the perfusion culture at 5 μl / min. Tubing was connected with a Luer-lock fitting from the IV media bag, through the peristaltic tubing, to the inlet of the perfusion chamber. Two 0.22 μm filters were placed in line with the flow before the hydrogel; one was placed just before the inlet and the other was placed at a Y-shaped stopcock between the media bag and the peristaltic pump. During long-term perfusion culture, media was added to the bag through this junction as needed. All air bubbles were expelled by pre-wetting the tubing all the way to the inlet leading to the perfusion chamber.

[0170] Epifluorescence imaging. Stitched images were acquired using a Nikon Eclipse Ti inverted epifluorescence microscope (Nikon Instruments Inc., Melville, NY) and a Zyla 4.2 sCMOS camera (Andor, South Windsor, CT). The degree of overlap was specified as 30%.

[0171] Fat pad implantation of angiogenic gel. All work involving animals was performed as described in a protocol approved by the Rice University IACUC. Male SCID / bg mice (6–10 weeks old, Charles River, strain code 250) received fat pad implantation of angiogenic gel. Briefly, mice were anesthetized with 2–4% isoflurane, and the abdomen was shaved and sterilized with betadine and isopropanol. Mice were administered 1 mg / kg Buprenorphine SR lab (Zoopharm) and 1 mL of 0.9% saline to prevent dehydration. An incision was made at the level of the linea alba using a scalpel and scissors, and the left and right gonadal fat pads were excised using Adson forceps. The angiogenic hydrogel was gently rinsed with PBS prior to implantation, and the gonadal fat pads were sutured to either side of the angiogenic hydrogel. Sterile 0.9% saline was used to gently wash the gel and maintain hydration during implantation. After suturing the fat pad, gel was introduced into the abdominal cavity and the muscles and skin were closed. After surgery, mice were monitored and care and husbandry was provided as needed.

[0172] Dextran injection. In some samples, fixable dextran 70 kDa conjugated to Texas Red (D1864, ThermoFisher) was prepared at 25 mg / mL in PBS. In other samples, fixable dextran 70 kDa conjugated to CF680 (Biotium, Cat. No. 80129) was prepared at 2.5 mg / mL in PBS. 15 min before euthanasia, 100 μL of dextran solution was injected via the tail vein. Gels were surgically explanted and photographed with a Sony A7R3 camera (Sony) and a Canon macro lens EF 100 mm 1:2.8L IS USM (Canon). Gels were then immersed in ice-cold 4% PFA and fixed at 4 °C for 48 h. After fixation, gels were washed three times with PBS for 20 min each and kept in PBS at 4 °C for imaging.

[0173] Islet isolation. Female Sprague-Dawley rats (251–275 g, Charles River, strain code 400) were utilized for islet isolation as described by a protocol approved by the Rice University IACUC. Rats were first humanely euthanized. Next, the abdomen was opened with a V-shaped cut, the diaphragm was cut, and the liver was inverted to expose the common bile duct. After injecting 15–20 mL of islet isolation buffer (buffer formulation) containing 1.5 mg / mL collagenase XI (C7657, Millipore Sigma) into the common bile duct, the pancreas was surgically removed and processed for enzymatic digestion of rat islets. Digestion included a 14-min incubation at 37°C, several washes with 10% BSA (50-253-923, Fisher Scientific), and gradient density separation (Products 99-690-CIS, 99-815-CIS, 99-691-CIS, 99-692-CIS, Corning). Islets were maintained in RPMI-1640 (10-040-CM, Corning) containing 10% FBS and 1% penicillin / streptomycin.

[0174] Induction of diabetes. Male SCID / bg mice (6–10 weeks old, Charles River, strain code 250) underwent induction of diabetes as described in an animal protocol approved by the Rice University IACUC. Mice were administered 70 mg / kg streptozotocin (572201, Millipore Sigma) dissolved in citrate buffer for 5 consecutive days.

[0175] Glucose monitoring. Blood glucose was quantified in a drop of blood produced by tail puncture using blood glucose strips and a blood glucose meter (BG1000, Clarity Diagnostics). Mice were considered diabetic if they had 1-h fasting glucose readings >400 mg / dL for 2 consecutive days. Mice were considered normoglycemic if they had non-fasting glucose readings ≤250 mg / dL.

[0176] Encapsulation of islets in alginate blocks. TMTD alginate and SLG100 (Pronova, Sandikva, Norway) were dissolved in two separate 0.8% saline solutions at 5% and 3%. These solutions were combined at a ratio of UPVLVG to SLG100 of 80:20 v / v. The alginate was sterile filtered through a 0.2 micron syringe filter and used to resuspend freshly isolated rat islets. Casting was performed by pipetting the alginate and islets into a mold of sterile 20 wt% PEGDA presoaked in crosslinking solution: 20 mM BaCl2 (Sigma-Aldrich 342920), 5 wt% mannitol (Thermo Fisher 33342). A thin polycarbonate membrane (Sterlitech 0.01 micron, 62x22mm, PCT00162X22100) was immersed in the crosslinking solution and folded over the side of the mold so that it covered the top and extended under the bottom. The edges of the rectangular membrane were pinched with tweezers to hold the membrane taut, and the mold was lifted and held in a bath of crosslinker (100mmx50mm crystallizing dish, PYREX) 2-3cm high. It was held in the crosslinker for 20 seconds, then removed from the bath and placed on a clean non-absorbent surface. The membrane was unwrapped and the mold was gently affixed onto a glass slide (Thermo Fisher, Cat. No. 12-544-7). The slide was then gently lowered back into the bath, securing the mold between the dish and slide. A weight (VWR 50mL conical tube filled with water) was then placed on top of the slide to press down on the mold, and crosslinking was continued for 1 hour. After 1 hour, the molds were removed from the bath and the blocks were gently peeled off. The blocks were then washed in a 50 mL conical containing 20 mL of HEPES buffer, adding 20 mL, draining 15 mL, and adding 15 mL of fresh buffer for each subsequent wash. The blocks were washed three times in the same manner with complete medium, then added to 10 mL of medium for storage in the incubator.

[0177] Islet viability. Live / dead staining solution was prepared by diluting calcein AM 1:200 v / v and ethidium homodimer-1 1:200 v / v from the Live / Dead kit (L3224, Thermo Fisher) in complete RMPI-1640. This staining solution was added 1:1 v / v to the medium of wells containing encapsulated islets. Encapsulated islets were kept at 37 °C and 5% CO2 for 30 min and then fixed with 4% PFA for 24 h at 4 °C. Stained and fixed islets were imaged using an EVOS FL (Thermo Fisher) and a Nikon Eclipse Ti epifluorescence microscope (Nikon).

[0178] Glucose-stimulated insulin release. Krebs buffer (KRB) was prepared as described in (Veiseh et al., Nature Medicine 22:306-311, 2016). KRB was prepared with 2 mM glucose (Sigma; G7528) (KR2) and 18 mM glucose (KR18). Encapsulated islets were incubated for 30 min in KR2 (wash), 1 h in KR2 (low), and then 1 h in KR18 (high). Supernatants were then collected from low (KR2) and high (KR20) wells, and insulin concentrations were quantified in duplicate by ultrasensitive rat insulin ELISA (ALPCO) without diluting the samples.

[0179] Statistical analysis. Statistical analysis was performed using GraphPad Prism 8 software (GraphPad Software). *P<0.05, **P<0.01, ***P<0.001, and ****P<0.0001.

[0180] Example 3 - Results To develop 3D printable and degradable copolymer gel substrates modified with immunomodulatory small molecules, we synthesized PEGDA, GelMA, and Met-RZA15 using established methods. LAP was prepared as previously described using equimolar amounts of dimethylphenylphosphinite and 2,3,6-trimethylbenzoyl chloride overnight at room temperature under argon (Figure 16A) (Fairbanks et al., 2009). In the next step, poly(ethylene glycol) diacrylates (PEGDA) of 3.4 kDa, 6 kDa, 10 kDa, 20 kDa, and 35 kDa were synthesized by reaction between poly(ethylene glycol) acryloyl chloride (in the presence of triethylamine overnight) (Figure 16A) (Miller et al., 2010). Gelatin methacrylate (GelMA) was synthesized by dropwise addition of methacrylic anhydride (0.1 mL per 1.0 g of gelatin) to a mixture of gelatin (type A, bloom 300) dissolved in 100 mM carbonate / bicarbonate buffer (pH = 9.5) at 50 C for 4 h (Figure 16B) (Shirahama et al., 2016). Met-RZA15, an immunomodulatory small molecule, was synthesized in a two-step method (Figure 16C). In the first step, RZA15 was synthesized by click chemistry reaction in the presence of copper iodide as a catalyst (Vegas et al., 2016b). In the next step, methacryloyl-modified RZA15 (Met-RZA15) was synthesized by the reaction between RZA15 and methacryloyl chloride in the presence of triethylamine under argon (Figure 16C). Intermediates and products were analyzed by ESI and / or NMR ( 1 H and 13 C) Analysis and characterization. Small molecule modified gels were generated by photopolymerization of 0.5 M aqueous Met-RZA15 into PEGDA (3.25 wt%), GelMA (10 wt%), and 17 mM LAP (Figure 16D). The batch sizes of these reagents are large enough to meet the large amounts of material required for these studies. [Table A]

[0181] The degree of acrylation of the polymers is consistent and comparable to that found throughout the literature. For bifunctional polymers such as PEGDA, where functional groups are present only at the end of each linear chain, a high degree of acrylation is necessary to fully incorporate each molecule within the hydrogel mesh. With >85% acrylation, the in-house synthesis is comparable to other expert chemists in the field and supports 3D printing of high-quality gels. For multifunctional polymers such as GelMA, where functional groups are present as amino acid side chains throughout the backbone, the desired degree of acrylation is application dependent. Increasing acrylation increases gel stiffness and improves handling; however, it reduces pore size, inhibits cell spreading, and slows collagenase degradation rates. For this reason, all studies were performed using GelMA with a moderate degree of acrylation (53.4 ± 8.3%). Acrylation was confirmed with 1H NMR by measuring the percentage of lysine residues in gelatin that have primary amine substitutions. With optimal acrylate degrees for large batches of biocompatible polymers (350 g PEGDA, 45 g GelMA), the synthesis pipeline supports the 3D printing of hundreds of angiogenic devices for parallel in vitro and in vivo studies. This scale has not been reported in the field before, increasing the likelihood of success for our group in pursuing vascular integration of transplanted islet structures. [Table B]

[0182] Analysis of polymer mechanics demonstrates low batch-to-batch variability. In addition to NMR confirmation of reproducible acrylation, reproducibility is further demonstrated by quantifying the ability of different polymer batches to behave consistently under shear deformation (storage modulus). This metric was obtained using an AR-G2 parallel plate rheometer. For each replicate, a 45 μL sample of polymer solution (containing 0.05 wt% Irgacure 2959) was added to the bottom plate and the top plate was lowered to give a gap size of 50 μm. The polymer solution was preconditioned for 30 seconds using an angular frequency of 5 rad / s and 10% strain, then 10 mW / cm of a 365 nm lamp was applied. 2 The change in storage modulus was monitored as the photocrosslinking reaction progressed, eventually resulting in a plateau value that could be compared to other samples.

[0183] With clinical targets in mind, implantable gel structures must be able to degrade in vivo. However, if these gels completely degrade before anastomosis occurs, exogenous islets will not be able to interact with the native circulation. A similar balance exists with regard to gel mechanics, where the gel must be stiff enough to be surgically manipulated. Here, we found that small (150 μL) 10% GelMA gels rapidly degrade in the presence of 0.2 mg / mL collagenase, and that the degradation rate can be tuned by incorporating PEGDA. By adding a relatively small amount of PEGDA to the gel formulation, collagenase took 4 times longer to degrade and shear stiffness increased 3-fold when measured using a parallel plate rheometer during UV exposure. These data suggest that the current formulation, 10% GelMA, 3.25% 3.4 kDa PEGDA, can meet the enzymatic and mechanical requirements (moderate degradation rate to support anastomosis, and sufficient stiffness to support surgical manipulation). Furthermore, the ability to tune these properties offers a unique opportunity to optimize angiogenesis in vivo without the need for exogenous stromal cells or growth factors. See Figures 17A-B.

[0184] We incorporate the anti-fibrotic molecule Met-RZA to prevent fibrosis in the 3D printed gels. To confirm the presence of Met-RZA on the surface of the gel, TOF-SIMS analysis was used to characterize the gel surface. Negative high-mass resolution depth profiles were obtained using a TOF-SIMS NCS instrument that combines a TOF-SIMS instrument (ION-TOF GmbH, Munster, Germany) with an in situ scanning probe microscope (NanoScan, Switzerland). Bunched 30 keV Bi 3+ ions (measurement current 0.2 pA) in a raster of 64x64 pixels using the primary probe for analysis (scan area 100 × 10 μm 2 ) was used. Adjustment of the charge effect was manipulated using a surface potential of -10 V. The cycle time was fixed at 200 μs (corresponding to a mass range of m / z = 0–3649 a.mu). In TOF-SIMS analysis, the gel surface was analyzed for the presence of small molecules (Met-RZA and RZA15) (Figure 18B–D). Negative polarity TOF-SIMS data show higher levels of Met-RZA (mass = 459.26, 1.6-fold) and fragmented RZA15 (mass = 390.98, 5.6-fold) in gels printed with Met-RZA small molecules compared to unprinted gels. This indicates the presence of immune-protective small molecules on the surface of 3D printed gels, which would reduce fibrosis of the implanted gels.

[0185] We demonstrate their ability to not only print patent vascular channels, but also to control circularity, which is important for endothelial coverage, proper hemodynamics, and leak-proof catheterization. This capability enables the realization of hydrogel vascular networks with complex 3D architectures, including various topologies of the vascular network. See Figure 19A-D.

[0186] Our demonstrated channel patency fabricated as designed indicates that we can successfully create a perfusable vascular network to support convection in a biocompatible angioplasty device. This also means that this 3D bioprinting technology can accommodate any vascular architecture, if the inventors deem different vascular designs more suitable for long-term survival of islets in the angioplasty device.

[0187] From the velocity profiles of both hydrogel architectures, the no-slip boundary condition appears to be a valid assumption since the velocity at either edge of the perfusable volume is near zero. The best-fit curves represent the data well, as indicated by the high r2 values. There are slight variations in the fluid dynamics of these gel architectures, but in general, flow appears unimpeded. Thus, these hydrogels would be able to accommodate blood flow if anastomoses were to occur in vivo and would support vascular incorporation to support islet cells. See Figures 21A-G.

[0188] Because we expect that these angiogenic devices will fully integrate into the host's circulatory system, it is important that the blood vessels do not rupture during the pressure changes of the cardiac cycle. Furthermore, the cyclic expansion of blood vessels provides mechanical cues to the underlying cell populations (e.g., endothelium, muscle, interstitium) that support healthy hemodynamics and promote homeostasis.

[0189] We demonstrate that these vascular networks can be fully endothelialized while retaining their lumen (Fig. 23C, inset cross section). Furthermore, there is evidence that HUVECs can actively degrade and invade the gel. We also observed HUVECs migrating in the parenchymal gel space, showing single cell migration (Fig. 23A, inset, arrows indicate a migrating single cell) and angiogenic sprouting (Fig. 23D). Endothelial cells are aligned in the direction of flow (Fig. 23B) and assume a classic cobblestone morphology (Fig. 23C, inset). Over the course of several days, HUVECs spread and form a confluent monolayer within a cylindrical lumen. These endothelialized gels can be stably maintained under perfusion for extended periods, and it is believed that this endothelial monolayer can be sustained until it reaches a quiescent state and is ready for in vivo implantation. The inventors believe that incorporation of an endothelialized network within the angioplasty device supports islet survival in vivo while promoting incorporation of the host vasculature for long-term survival and functionality.

[0190] The high viability of endothelial cells after several days of perfusion suggests that the endothelialized vascular network remains robust for extended periods of time and further suggests that the endothelial monolayer has matured. The endothelial monolayer can be maintained in this quiescent state for several days without losing viability. This allows grafts to be prepared in advance and seeded when donor islets become available.

[0191] After successfully seeding the 3D printed gel with endothelial cells, islets were introduced into the gel by pouring alginate (Figure 25A). A pouring method was developed to ensure that the cross-linked alginate surrounded the anchors (Figure 25B), holding the alginate and islets in place. In a sterile environment, two types of alginate, UP VLVG modified with RZA15 and SLG100 (NovoMatrix, Sandvika, Norway), were mixed in a 60:40 ratio and filtered to produce TMTD alginate. 60 μl of TMTD alginate was pipetted into the gel chamber and cross-linked with 100 mM CaCl2 and 1 mM BaCl2 solution for 30 minutes. This process was repeated to fill the chamber with alginate and islets. Successful cross-linking was confirmed by vibratome cross-sections showing solid alginate throughout the block. By following the above procedure, TMTD cast gels can be consistently produced. This alginate casting procedure results in a structure that can be firmly handled without falling apart. Successful alginate casting allows for the encapsulation of pancreatic islets.

[0192] TMTD alginate was prepared and mixed with human pancreatic islets. The gel containing alginate and islets was incubated with Live / Dead stain (1:200 calcein AM, 1:200 ethidium homodimer in complete CMRL-1066) for 30 minutes. The gel containing islets was then fixed and imaged with a Nikon Ti Eclipse epifluorescence microscope.

[0193] For GSIS, 3D-printed angiogenesis devices were infused with TMTD alginate and human islets. The 3D-printed angiogenesis devices were incubated in 2 mM Krebs buffer for 30 min (prewash), 60 min (wash), and 60 min (low glucose). They were then incubated in 18 mM Krebs buffer for 60 min (high glucose). Low and high glucose supernatants were analyzed in duplicate by ultrasensitive human ELISA (Alpco, Salem, NH). Stimulation index was calculated as the ratio of insulin produced in high glucose divided by insulin produced in low glucose.

[0194] The demonstration of >90% viability and a stimulation index >2 indicates that our device and method of islet encapsulation is compatible with fragile islet biology, enhancing the translational value of this approach for allogeneic and xenogeneic islet transplantation.

[0195] We implanted gels endothelialized with GFP-HUVECs to evaluate the in vivo anastomosis of gels with patterned vasculature (Figure 26A). Gels were seeded with GFP-HUVECs and maintained in perfusion culture for 6 days (Figure 26B). To promote vascular anastomosis, highly vascularized gonadal fat pads were wrapped around the gels and sutured to the gels (Figure 26C). Prior to euthanasia, mice were administered human-specific and mouse-specific lectins via tail vein injection. Gels were explanted at 3 weeks, fixed, and maintained at 4C for analysis. The 3D printed channels retained their snake-like topology and remained patent during implantation in vivo. This indicates that the vasculature uniquely patterned in vitro translates to a similar pattern in vivo. Visualization of blood in the gel channels (Figure 26D) is a promising sign of vascular incorporation with the gels. Closely associated blood vessels (Figure 26D) indicate that the gel is degradable and capable of vascular integration with the host. These findings encourage the potential of this approach to yield patterned vasculature in vivo for encapsulated islet transplantation.

[0196] Example 4 Development of patterned multilayer hydrogels with various biomaterial properties. Patterned hydrogels are developed as follows: Multiple hydrogel formulations can be utilized to create multilayer patterned hydrogels with various biomaterial properties such as degradability, stiffness, cellular components, and cell adhesion properties.

[0197] Three-layer hydrogels with a non-degradable middle layer along with degradable bottom and top layers can be developed as follows: hyaluronic acid modified with norbornene functional groups (norHA), combined with a photoinitiator (e.g., LAP), and a degradable peptide crosslinker (e.g., CGPQGIWGQGCR (SEQ ID NO: 1)), result in patterned fully synthetic hydrogel formulations with degradable peptide crosslinks, allowing for the specification of degradation profiles.

[0198] After the non-degradable portion is printed, for example using a stereolithography approach, the printing solution can be exchanged for a non-degradable one, for example 6.0 kDa PEG-diacrylate, and the patterned components of the non-degradable PEG-diacrylate can be printed with the same device.

[0199] Finally, the non-degradable gel formulation can be exchanged for a degradable gel formulation, e.g., norHA+LAP+CGPQGIWGQGCR (SEQ ID NO: 1) peptide crosslink, allowing the creation of multi-layered biomaterials with different biomaterial properties.

[0200] Hyaluronic acid was functionalized with norbornene to generate norbornene-modified hyaluronic acid (NorHA) by mixing norbornene and hyaluronic acid in the presence of Boc2O, DMAP, and DMSO and stirring at 45 C for 20 h. Hyaluronic acid gels were printed with pendant cell adhesive peptides (exposure to blue 405 nm light) with lithium acylphosphinate (LAP) and crosslinking of degradable peptides by addition of cell adhesive thiol-terminated peptide (CGRGDS (SEQ ID NO:2)) and MMP-sensitive dithiol crosslinking peptide (CGPQGIWGQGCR (SEQ ID NO:1)). The peptide sequences CGPQGIWGQGCR (SEQ ID NO:1) (highly degradable, HD, 1261.42 g / mol) and CGRGDS (SEQ ID NO:2) (thiol-terminated cell adhesive peptide) were custom synthesized by Aapptec (Louisville, KY). Peptides were supplied as trifluoroacetate salts with a purity of >95%. Peptides were deaerated and stored at -80°C under argon (to minimize disulfide formation) until needed (Miller et al., Biomaterials, 31(13:3736-43, 2010).

[0201] The gel was purified as follows: NorHA and 6 kDa PEG-diacrylate were dissolved in PBS at 4% by weight each. The thiol-terminated degradable cross-linking peptide CGPQGIWGQGCR (SEQ ID NO: 1) was added at 10 mM. LAP was added at 34 mM. The light absorber tartrazine was added at 2.25 mM.

[0202] All methods disclosed and claimed herein can be made and performed without undue experimentation in light of this disclosure. Although the compositions and methods of the present invention have been described in terms of preferred embodiments, it will be apparent to those skilled in the art that modifications may be applied to the methods and steps or sequence of steps of the methods described herein without departing from the concept, spirit and scope of the present invention. More specifically, it will be apparent that certain agents that are both chemically and physiologically related may be substituted for the agents described herein while the same or similar results would be achieved. All such similar substitutes and modifications apparent to those skilled in the art are deemed to be within the spirit, scope and concept of the present invention as defined by the appended claims.

[0203] IX. References The following references, to the extent that they provide exemplary procedural or other details supplementary to those set forth herein, are specifically incorporated herein by reference. TIFF0007610517000003.tif208160TIFF0007610517000004.tif246160TIFF0007610517000005.tif238160TIFF00076105170 00006.tif238160TIFF0007610517000007.tif225160TIFF0007610517000008.tif237161TIFF0007610517000009.tif141160

Claims

1. a degradable outer housing comprising a degradable polymer and an inner non-degradable portion comprising a non-degradable polymer; An implantable device comprising: the degradable outer housing and / or the inner non-degradable portion include a patterned architecture that promotes vascularization of the implantable device; the inner non-degradable portion comprises a natural or synthetic hydrogel, the hydrogel comprising greater than 80% water; The degradable outer housing and / or the inner non-degradable portion are loaded with cells or drugs. The implantable device.

2. The implantable device of claim 1 , wherein the synthetic hydrogel is poly(ethylene glycol) diacrylate (PEGDA).

3. 3. The implantable device of claim 1 or 2, further comprising gel methacrylate (GelMA).

4. The implantable device of any one of claims 1 to 3, wherein the hydrogel comprises 10% GelMA and 3-4% PEGDA having a molecular weight of 3.4 kDa.

5. The hydrogel is an acrylate-(PEG-protein) ｎ 2. The implantable device of claim 1, comprising monomers of -PEG-acrylate, optionally where PEG is 2-35 kDa and the protein is greater than 100 Da with a total MW of at least 500 kDa or n=2-500 repeat units.

6. 10. The implantable device of claim 1, wherein the hydrogel comprises acrylamide, vinyl sulfone, norbornene-substituted acrylate groups, alginate, silk, dextran, chondroitin sulfate, hyaluronic acid, cellulose, heparin, poly(caprolactone), or multi-component versions thereof.

7. 2. The implantable device of claim 1, wherein the cells comprise cells from heterologous tissue, cells from a cadaver, stem cells, cells derived from stem cells, cells from a cell line, primary cells, reprogrammed cells, reprogrammed stem cells, cells derived from reprogrammed stem cells, genetically engineered cells, or combinations thereof.

8. The implantable device of claim 1 , wherein the cells are human cells.

9. The implantable device of claim 1 , wherein the cells are insulin-producing cells.

10. The implantable device of claim 1 , wherein the cells are pancreatic islet cells.

11. The implantable device of any one of claims 1 to 10, wherein the degradable outer housing or the inner non-degradable portion is loaded with an oxygen-releasing compound.

12. The implantable device of any one of claims 1 to 11, wherein the degradable outer housing or the inner non-degradable portion is loaded with a drug or angiogenic cells.

13. The implantable device of any one of claims 1 to 6, wherein the inner non-degradable portion is loaded with a biosensor.

14. 14. The implantable device of claim 13, wherein the biosensor is capable of measuring blood oxygen, blood glucose, tumor antigens, cholesterol, creatinine, hemoglobin A1C, insulin, glucagon, hormone levels, levels of small molecule drugs, DNA, RNA, cytokines, or circulating tumor cells.

15. The implantable device of any one of claims 1 to 14, wherein the inner non-degradable portion is patterned to allow for vascularization once implanted.

16. The implantable device of any one of claims 1 to 15, wherein the patterned portion comprises a vascular geometry.

17. The implantable device of any one of claims 1 to 16, wherein the device comprises multiple layers of hydrogel.

18. An implantable device according to any one of claims 1 to 17 for monitoring a state or condition of a subject, the device being adapted to be implanted in said subject, the device comprising a sensor or biosensor.

19. 18. An implantable device according to any one of claims 1 to 17 for treating a medical condition or disease in a subject, the device being adapted to be implanted into the subject, the device comprising a therapeutic agent.

20. An implantable device according to any one of claims 1 to 17 for creating an angiogenic device in a subject, the device being used to be implanted in said subject.

21. An implantable device according to any one of claims 1 to 17 for creating angiogenic tissue around an implant or prosthesis in a subject, the device being adapted to be implanted into the subject, the implant or prosthesis being a neural interface device.

22. An implantable device according to any one of claims 1 to 17 for creating tissue augmentation around an implant or prosthesis in a subject, the device being adapted to be implanted in the subject, the implant or prosthesis being a breast implant or tissue filler.

23. The implantable device of any one of claims 1 to 17, wherein the degradable polymer comprises a patterned vascular network embedded therein.

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