science.sciencemag.org/content/365/6452/482/suppl/DC1

Supplementary Materials for

3D bioprinting of collagen to rebuild components of the human heart

A. Lee, A. R. Hudson, D. J. Shiwarski, J. W. Tashman, T. J. Hinton, S. Yerneni, J. M. Bliley, P. G. Campbell, A. W. Feinberg*

*Corresponding author. Email: feinberg@andrew.cmu.edu

Published 2 August 2019, Science 365, 482 (2019) DOI: 10.1126/science.aav9051

This PDF file includes: Materials and Methods Figs. S1 to S15 Table S1 Captions for movies S1 to S10 References Other supplementary material for this manuscript includes: Movies S1 to S10

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MATERIALS AND METHODS 3D Bioprinter Setup All FRESH printing was performed on custom-built 3D bioprinters based on open-source modifications of consumer-grade desktop 3D printers, similar to our previously published designs (14, 16). The thermoplastic extruder on a range of 3D printers (e.g., PrintrBot Simple Metal, MakerBot Replicator 2X, Flashforge Creator Pro) was replaced with our open-source Replistruder 3 syringe pump extruder (fig. S1) released under the Creative Commons CC-BY-SA license and available for download as STL files on the NIH 3D Print exchange website (https://3dprint.nih.gov/discover/3DPX-009853). The Replistruder 3 improves upon previously published Replistruder versions with a simpler design to 3D print, increased rigidity for precision bioink extrusion and a modular architecture for rapidly changing syringes (fig. S1A). For single-material printing, the Replistruder 3 was mounted on a custom-designed carriage and loaded with a Hamilton gas tight syringe with stainless steel needle (fig. S1B). For multi-material printing, two or three Replistruder 3 syringe pumps were mounted on a custom-designed carriage for the printer (fig. S1C to S1F). Needles on each syringe were first adjusted in the x- and y-axis via set screws mounted to the syringe carriage in alignment to a reference grid laser-etched onto the build plate (fig. S1E). Needles were then adjusted in the z-axis using set screws on the motor mount until they were aligned to the same z height. The known distance between the needle tips was then entered into the Slic3r slicing software (version 1.2.9, https://slic3r.org/), and a tool change G-code was inserted to switch between the two syringes. Two petri dishes filled with deionized (DI) water were placed on each side of the print container to prevent clogging of the needles not actively printing. FRESH Support Bath Generation and Characterization FRESH v1.0 support bath was prepared as previously described using a mechanical blending method to generate gelatin microparticles (16). FRESH v2.0 support bath was prepared using a complex coacervation method to produce gelatin microparticles with smaller and more uniform shape and size. First, 2.0% (w/v) gelatin Type B (Fisher Chemical), 0.25% (w/v) Pluronic® F-127 (Sigma-Aldrich) and 0.1% (w/v) gum arabic (Sigma-Aldrich) were dissolved in a 50% (v/v) ethanol solution at 45ºC in a 1 L beaker and adjusted to 6.25 pH by addition of 1M hydrochloric acid (HCl). To form the slurry of gelatin microparticles, the beaker was then placed under an overhead stirrer (IKA, Model RW20), sealed with parafilm to minimize evaporation, and allowed to cool to room temperature while stirring overnight. The resulting slurry was then divided into 50 mL conical tubes and centrifuged at 300 g for 5 min to compact the gelatin microparticles. The supernatant was then removed and gelatin microparticles were resuspended in a washing solution to remove the ethanol and Pluronic® F-127. The gelatin slurry was then washed three times with washing solution, which depended on the bioink to be used, at 1000 g for 2 min. Slurry to be used with alginate bioink required a washing solution of 0.1% (w/v) CaCl2 (Sigma-Aldrich), while slurry to be used with acidified collagen bioink typically used a washing solution of 50 mM 4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid (HEPES) (Corning), pH 7.4, or 1X phosphate buffered saline (PBS), pH 7.4. The gelatin slurry in its uncompacted state in the washing solution was stored at 4ºC for up to 1 month. Prior to printing, the uncompacted slurry was degassed in a vacuum chamber for 15 min, followed by centrifugation at 2000 g for 5 min. The supernatant was removed, and the gelatin slurry transferred into a print container of appropriate volume, typically 50% larger than the construct to be printed. For sterile support preparation, all hardware was either

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autoclaved or cleaned with 70% ethanol and then sterilized with ultraviolet (UV) in a biosafety cabinet. Coacervate materials were sterilized through sterile filtering or exposure to UV light in a biosafety cabinet, with the entire coacervation and washing steps occurring in a biosafety cabinet. To assess microparticle size, uniformity and distribution, compacted gelatin slurry was diluted in washing solution and stained with black food coloring (McCormick & Co.). The microparticles were imaged in brightfield using an inverted microscope (TS100, Nikon) with a digital camera (D7000 SLR, Nikon) and a macro was created in ImageJ (National Institutes of Health) to analyze microparticle Feret diameter. To measure the rheological properties of the gelatin support bath, compacted slurry prepared for printing was loaded onto a rheometer (Gemini 200 Rheometer, Malvern) equipped with a 40 mm diameter, 4° cone with a 150 μm separation at 25ºC. A strain-controlled amplitude sweep (fig. S2A) from 0.1% to 100% strain at a frequency of 1 Hz was first used to determine the linear viscoelastic region and apparent yield stress of the support. Afterwards, a frequency sweep (Fig. 1G) was performed from 0.1 to 150 Hz while the storage (G’) and loss moduli (G”) were recorded in Microsoft Excel and plotted with GraphPad Prism 7.0. Bioink Preparation A total of five different bioinks were used in these studies. Alginate bioink was prepared by making a 4% (w/v) solution of sodium alginate (FMC BioPolymer) in deionized (DI) water with 0.1% Alcian blue (J60122, Alfa Aesar) dye added to aid visualization. Collagen bioink was prepared by diluting sterile 35 mg/mL collagen type I (LifeInk200, Advanced Biomatrix) with 0.24M acetic acid (VWR) in a 2:1 volume ratio by extruding back and forth between two mated syringes 40 times to ensure thorough mixing. The acidified collagen bioink was centrifuged at 3000 g for 5 min to remove bubbles and then transferred to gastight syringes of volume 1, 2.5 or 10 mL for printing (Gastight Syringe, Hamilton Company). A 24 mg/mL acidified collagen bioink was used for the majority of the prints, including collagen strands, multi-material prints, perfusion tubes, ventricle prints, whole heart prints, multiscale vasculature prints, and valve prints. For the in vivo studies, 12 mg/mL acidified collagen bioink was used as is or modified with 100 ng/mL vascular endothelial growth factor (VEGF) and 60 µg/mL fibronectin. For micro-computed tomography (µCT) imaging, 24 mg/mL acidified collagen bioink was modified with 2.0% (w/v) barium sulfate powder (Sigma-Aldrich). Methacrylated hyaluronic acid (MeHA) bioink was prepared by making a 1.5% (w/v) solution of MeHA (PhotoHA, Advanced Biomatrix) in 1X PBS at 4ºC for 1 hour. A 10% (w/v) stock solution of Photoinitiator (Irgacure, Advanced Biomatrix) was dissolved in methanol. Enough photoinitiator stock was then added to the solution of MeHA to produce a final concentration of 0.1% (w/v) photoinitiator. The solution was covered and stirred at 4ºC overnight. Fibrinogen bioink was prepared by mixing fibrinogen (MP Biomedicals) and xanthan gum (MP Biomedicals) in 1X PBS to produce a final concentration of 40 mg/mL fibrinogen and 0.2% (w/v) xanthan gum. For the preparation of the cell bioink, human embryonic stem cell (hESC)-derived cardiomyocytes were passaged and resuspended in CDM3 (chemically defined medium, 3 components) medium, filtered through 40 µm nylon cell strainer (STEMCELL technologies) to remove large cell aggregates, centrifuged 200 g for 5 min and resuspended in 20 mg/mL fibrinogen and CDM3 medium, and transferred into sterile, capped, 10 mL plastic syringes (BD). Syringes containing the

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cell suspension were fitted into custom centrifuge adapters and pelleted again inside the syringe at 200 g for 7 min. Supernatant was removed from the syringe, and the cell pellet transferred to a sterile gastight glass syringe via a Luer lock mate adapter. Gentle mixing between the two syringes was carried out to disperse the condensed cell pellet to form a uniform cell bioink. Final concentration of cell bioink was approximately 300 million cells/mL. All syringes, needles, adapters, and syringe accessories were sterilized in the autoclave. FRESH Printing of Collagen Type I 3D bioprinting of collagen type I was performed using the FRESH v2.0 support bath and our custom-designed, open-source 3D bioprinters (14, 16). All digital models were created in computer-aided design (CAD) software (SOLIDWORKS 2016, Dassault Systèmes) or downloaded from the BodyParts3D database (http://lifesciencedb.jp/bp3d/) (24) or the National Institutes of Health (NIH) 3D print exchange (https://3dprint.nih.gov/discover/3dpx-000452). Minor repairs and modifications to surface details of the downloaded digital models were made in Meshmixer (Autodesk) to improve print fidelity. All 3D models were then exported as an STL format file and processed in Slic3r slicer software to generate G-code instructions for the printer. In general, slicer settings were a speed of 23 mm/s, layer height of 40-100 μm, 0-2 perimeters, and 30-80% rectilinear infill. Prior to printing, collagen bioink was transferred into a gastight glass syringe (10 mL for the organ-scale heart prints, 2.5 mL for all others) and mounted into a Replistruder 3. A needle was fitted to the syringe and primed. The needle was typically stainless steel 0.5” in length with an inner diameter (ID) of 80-250 µm (Jensen Global), with most printing performed with a 150 µm ID needle. We also used a 10 µm ID pre-pulled glass pipette tip with Luer fitting (World Precision Instruments) to determine maximum resolution, and for the whole heart print we used a 0.5” long, 200 µm ID needle epoxied in to a 2” long 840 µm ID needle that served to both extend and reinforce the deposition needle from deflection. The Replistruder 3 with syringe was mounted onto the 3D bioprinter using the custom-designed extruder carriage. A container large enough to hold the construct to be printed was filled with FRESH support bath and secured to the print platform with a thin layer of vacuum grease. The needle was manually positioned in the xy center of the container and lowered to 1 mm above the bottom of the print container. The 3D print was then started using the open-source printer control software Pronterface (version 1.6.0, http://www.pronterface.com). All collagen constructs were printed at room temperature (22ºC), with sterile printing performed inside a biosafety cabinet. Upon print completion, the container was removed from the platform and incubated at 37ºC to melt the FRESH support bath and release the printed construct. Once released, collagen constructs were washed with 50 mM HEPES, pH 7.4 solution, unless otherwise specified, to remove the melted gelatin support, and then used for specific applications. Optical Imaging of FRESH Printed Collagen Constructs To evaluate printed filament structure and microscale print quality we performed multiple types of optical imaging. Printed collagen filaments were imaged using phase-contrast imaging on an inverted microscope (TS100, Nikon) with a digital camera (D7000 SLR, Nikon). Printed collagen constructs at the millimeter scale and above were imaged using a digital SLR camera (Model 70D, Canon) under oblique and general illumination. Microscale 3D imaging of collagen filaments and constructs was performed using confocal reflectance and second harmonic generation imaging on a Leica SP5 multiphoton microscope equipped with a 25X 0.95 NA lens, 405, 488, 515, 561, 647 visible laser lines, and an InSight X3 multiphoton laser (Spectra-Physics). Reflectance microscopy

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was performed with 488 nm excitation followed by gating fluorescence collection on the spectral detector to 488 nm +/- 20 nm. Nyquist optimized confocal sectioning was performed to acquire 3D reconstructions of the printed constructs. The second harmonic generation for collagen imaging was performed using 830 nm excitation with the InSight X3 laser and a 491 nm dichroic mirror for collection with Leica HyD detector. Visualization and surface rendering of 3D imaging data was done in Imaris 9.1 (Bitplane). Printing four unique bioinks with distinct cross-linking mechanisms Four bioinks (alginate, acidified collagen, fibronectin, MeHA) were each prepared separately as previously described. To label each bioink with a unique fluorescent color, human fibronectin (Corning) was conjugated to Alexa-Fluor 405, 488, 555, and 633 NHS Esters (Thermo Fisher Scientific) and added to the alginate, MeHA, acidified collagen, and fibrinogen, respectively. To fluorescently label the fibronectin, succinimidyl ester (SE) buffer was prepared by making a 0.84% (w/v) solution of sodium bicarbonate (Fisher Scientific) in DI water. Afterwards, 200 µL SE buffer was combined with 400 µL fibronectin in a 1:2 volume ratio. All Alexa-Fluor dyes were first dissolved in 100 µL sterile dimethyl sulfoxide (DMSO) (VWR). After gently mixing 10 µL of Alexa-Fluor dye with fibronectin solution, samples were let to sit at room temperature, covered, for 1 hour. Each sample of dyed fibronectin was then dialyzed against SE buffer at 4ºC for 2 hours to remove excess dye. SE buffer was then replaced with 1X PBS and let to dialyze at 4ºC for 2 hours. The 1X PBS was then replaced with DI water and dialyzed at 4ºC overnight. The following morning, the labeled fibronectin solutions were aliquoted and stored at -20ºC. To fluorescently label each bioink, stained fibronectin solution was added to each bioink in a 1:10 volume ratio. All inks were centrifuged at 4000 g for 2 mins to remove air bubbles. The FRESH v2.0 gelatin support bath was washed with 0.1% (w/v) CaCl2, 50 mM HEPES, 1 U/mL thrombin to support the cross-linking of alginate, acidified collagen and fibronectin solution, respectively, and compacted at 2000 g for 5 minutes. To demonstrate filament resolution, a square lattice was printed from each ink into separate containers using the support bath described above. A 10x10x5 mm square was designed in Fusion 360 (Autodesk) and exported as an STL file. The bath supporting the MeHA ink was exposed to UV light (OmniCure S2000, Excelitas Technologies) during printing as well as an additional 5 minutes each on the top and bottom sides of the print container to increase crosslinking. After printing, all inks were incubated at 37ºC for 1 hour to remove the support bath. Each lattice was imaged using a calibrated overhead digital camera (D7000 SLR, Nikon). To demonstrate the ability of the bath to print all four bioinks together, the letters “CMU” embedded in an 8x18x2 mm frame was exported as an STL file using Fusion 360. Using the same support bath described above, the CMU frame was then printed using acidified collagen and each individual letter of “CMU” was then printed with a different bioink. The letter “C” was printed from MeHA with crosslinking by exposing the bath to UV light during printing and an additional 5 minutes each on the top and bottom of the print container. The letters “M” and “U” were printed from labeled fibrinogen and alginate, respectively. After printing, the gelatin support bath was removed by incubating at 37ºC for 1 hour. The print was then imaged on a LSM 700 laser scanning confocal microscope (Zeiss) with a 10X 0.4 NA objective.

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Primary and Cell Line Culture and Cardiomyocyte Differentiation Viability of cells within perfused 3D constructs was evaluated using murine C2C12 myoblasts (CRL-1772, ATCC). The C2C12s were cultured at 37°C under 5% CO2 with Dulbecco’s modified Eagle’s medium (DMEM) (15-013-CM, Corning) supplemented with 10% (v/v) Fetal Bovine Serum (FBS) (89510-186, VWR), 1% (v/v) L-glutamine (25030-081, Life Technologies), and 1% (v/v) penicillin-streptomycin (15140-122, Life Technologies). Media was exchanged every 2 days and C2C12s were passaged prior to reaching 80% confluence. Human cardiomyocytes for the 3D bioprinted ventricles were differentiated from HES3 hESCs based on minor modifications of established protocols and media formulations (Table S1) (25). Briefly, hESCs were expanded in Essential 8 (E8) media (A1517001, Life Technologies) on 6 well plates coated with 12 µg/cm2 Geltrex (A1413301, Life Technologies) (26). Media was changed daily and hESCs were passaged every four days at 1.25 X 105 cells/well in E8 supplemented with 2 µM thiazovivin (S1459, Selleck Chemicals) to enhance survival during passaging. To obtain an optimal density for cardiomyocyte differentiation, HES3 hESCs were detached from the plate with TrypLE express (12604013, Thermo Fisher Scientific) and seeded at 16,000 cells/cm2 in Geltrex-coated dishes with E8 with 2 µM thiazovivin. The hESCs were then allowed to proliferate until ~50% confluence was achieved. Cardiomyocyte differentiation was then carried out using a combination of previously described protocols (25, 26). On day 0, the media was changed to RPMI/B27 with 6 µM CHIR99021 (C-6556, LC laboratories). On day 2, the media was changed to RPMI/B27 with 2 µM Wnt-C59 (S7037, Selleck Chemicals). On day 4 and 6, media was changed to RPMI/B27. On day 8 and 10, media was changed to CDM3. On day 12, spontaneously beating cardiomyocytes were passaged for metabolic purification using a lactate-supplemented media (CDM3L). Specifically, beating cardiomyocytes were washed with 1X PBS and detached from the surface with TrypLE express for 15 min at 37℃. Detached cells were pipetted into DMEM/F12 media and centrifuged at 200 g for 7 min to pellet the cells. cardiomyocytes were seeded on 12 µg/cm2 Matrigel (356231, Corning)-coated plates with CDM3L and purified for 5 days. After purification, cardiomyocytes were maintained in CDM3 for up to 28 days prior to FRESH printing experiments. Cell adhesion and spreading on the surface of 3D bioprinted collagen constructs, specifically heart valves leaflets, was evaluated using human umbilical vein endothelial cells (HUVECs, CC-2519, Lonza). The HUVECs were cultured in endothelial cell growth media (CC-3124, Lonza) supplemented with penicillin (100 U/ml) and streptomycin (100 mg/ml) at 37°C under 5% CO2. Media was exchanged every 2 days and HUVECs were passaged prior to reaching 80% confluence. Millimeter-Scale Collagen Tubes – Fabrication, Perfusion and Characterization Perfusable collagen tubes were designed to approximately the scale of small arteries, comparable to regions of the left anterior descending artery in the human heart. The specific CAD model (fig. S6A) consisted of a tube with 1.4 mm ID lumen, 0.3 mm wall thickness and 9 mm freestanding length. Each end of the tube was embedded within a block-like region of collagen to provide mechanical support for the freely-suspended region of the tube and for needle insertion during perfusion. The central region of the construct consisted of a tissue casting chamber with a volume of 600 µL, designed to form a tissue that was diffusion-limited to assess cell viability and tissue compaction during perfusion. The construct was printed with the tube oriented within the xy build

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plane, and the tube walls were printed as perimeters at a speed of 10 mm/s to ensure high print fidelity and lumen patency. Once printed, perfusion tube constructs were released from the FRESH support by heating to 37ºC, washed gently with 50 mM HEPES and sterilized under UV (PSD Pro Series, Novascan) for 15 min prior to use. Fidelity, patency and wall integrity of the printed collagen tube was assessed by perfusion using fluorescently-labeled dextran. Perfusion within the tube lumen and diffusion through tube wall was visualized using 10 kDa rhodamine dextran (D1A24, ThermoFisher) and 70 kDa fluorescein dextran (D1A23, ThermoFisher) flowed sequentially through the tube using a peristaltic pump (ISM934C, Ismatec). Epifluorescent time-lapse imaging was performed on a stereomicroscope (SMZ1500, Nikon) with a CMOS camera (Prime 95B, Photometrics) and X-cite lamp (Excelitas Technologies). A kymograph was constructed from the time series in ImageJ and used to quantify relative dextran diffusion through the tube wall and in to the surrounding tissue casting chamber. Long-term assessment of cellularized collagen tube constructs was performed using custom-designed perfusion chambers under sterile culture conditions. Perfusion chambers (fig. S5A) designed in SOLIDWORKS to hold the collagen construct and interface it with inlet and outlet needles were fabricated using a multi-step process. First, the main chamber was printed with Dental SG resin (RS-F2-DGOR-01, Formlabs) on a Form 2 SLA printer (Formlabs) (fig. S5B). A polylactic acid (PLA) insert was 3D printed with a Replicator 2X (MakerBot) and placed on top of the perfusion chamber to form a temporary mold around which polydimethylsiloxane (PDMS) Sylgard 184 (Dow Corning) was cast and cured to form a rectangular well (fig. S5C). Stainless steel, 310 µm ID, 1” long deposition needles (JG24-1.0HP, Jensen Global) were inserted into the side of the perfusion chamber and central PLA insert during PDMS curing to form through holes within the PDMS well for future perfusion needle placement (fig. S5D). Prior to use, perfusion chambers were sterilized with an autoclave. For perfusion studies, collagen tube constructs were carefully transferred into the PDMS wells in the perfusion chambers and stainless-steel needles were inserted for inflow and outflow (fig. S5E). C2C12s in a 3D hydrogel were cast around the collagen tube within the tissue casting chamber, consisting of 2.5 mg/mL rat tail collagen type I (354249, Corning), 1.7 mg/mL Matrigel (354234, Corning), 10% (v/v) 10X PBS, 2.3% (v/v) 1N NaOH, and 37.5 X 106 C2C12 cells/mL. A 600 µL volume of this cell and collagen mixture was pipetted around each collagen tube and gelled in the incubator at 37ºC for 1 hour. Collagen tube constructs with cast C2C12s were then cultured as static controls (N=3) or perfused (N=3) within the incubator. The 1/16” (1.6 mm) ID silicone tubing (EW-95802-02, Cole Parmer), male and female Luer connectors (51525K291 and 51525K141, McMaster Carr), perfusion needles (JG24-1.0HP, Jensen Global), and media bags (32C, St. Gobain) were autoclaved and then connected into a flow loop (fig. S5, F and G). Collagen tube constructs were perfused at a flow rate of 0.4 mL/min using a peristaltic pump (ISM934C, Ismatec) for 5 days at 5% CO2 and 37ºC using DMEM (15-013-CM, Corning) with 10% (v/v) FBS (89510-186, VWR), 1% (v/v) L-glutamine (25030-081, Life Technologies), and 1% (v/v) penicillin-streptomycin (15140-122, Life Technologies). The non-perfused controls were cultured for 5 days in the perfusion chambers with 3 mL of medium on top of the construct for 5 days with daily media exchanges.

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To determine C2C12 viability, constructs were removed from the perfusion chambers at 5 days and rinsed with Tyrode’s solution (T2145, Sigma-Aldrich) for 10 min on an orbital shaker. Constructs were then cross-sectioned with a scalpel and stained with 2 µM calcein AM and 4 µM ethidium homodimer (L3224, Life Technologies) and 4′,6-diamidino-2-phenylindole (DAPI) for nuclear staining according to the manufacturer’s protocol. Large area tile-scan images of the construct cross-section were performed using a Zeiss LSM 700 laser scanning confocal microscope with a 10X objective. Percent viability was determined with Imaris image analysis software using “spot” detection, where percent viability was equal to number of live cells (calcein-positive) divided by the total number of nuclei (DAPI). Percent viability was determined as a function of distance from the top surface of the construct (i.e. not touching the sides of the PDMS well in the perfusion chamber). A depth of 600 µm from the top surface of the construct was analyzed because the perfused constructs compacted much more than non-perfused controls. In Vivo Assessment of FRESH v2.0 Printed Collagen Scaffolds In vivo studies were performed with collagen disk-shaped scaffolds 10 mm in diameter and 5 mm in height and were FRESH v2.0 printed or punched out from cast sheets, as controls. Collagen disks were printed under sterile conditions with 100 µm layer height, 85% rectilinear infill, and 0 perimeters to achieve maximum collagen content. Control collagen disks were cast in 48-well plates and cut to size using a 1 cm biopsy punch. C57BL/6 male mice (6 to 8 weeks old) were utilized in this study and animal care and experimental procedures were carried out at Carnegie Mellon University in accordance with the NIH Guide for the Care and Use of Laboratory Animals under an approved Institutional Animal Care and Use Committee (IACUC) protocol. Two sets of experiments were performed (N = 6 mice per experimental condition). First, disks printed or cast using collagen at 12 mg/mL were implanted for 3, 7, and 14 days (N = 6 for all conditions at each time point) to evaluate the time course of cellular infiltration and scaffold remodeling as a function of fabrication method. Second, disks printed (N = 6) or cast (N = 6) using collagen at 12 mg/mL with the addition of 100 ng/mL VEGF (100-20, PREPROTECH) and 50 µg/mL fibronectin (F2006, Sigma-Aldrich) were implanted for 10 days to evaluate cellular infiltration and host vascularization. Scaffolds were implanted in vivo based on established methods (27). Scaffolds were implanted subcutaneously on either side of the dorsum under general anesthesia (2% isoflurane). For the vascularization experiments, prior to sacrificing the mice, 100 µL of DyLight®

488 labeled Lycopersicon esculentum lectin (DL-1174, Vector Laboratories) was injected via the tail vein. Upon removal, scaffolds were cleaned of connective and adipose tissue, and then fixed in 10% formalin overnight before tissue processing and embedding in paraffin. A combination of imaging techniques was used to assess remodeling of the collagen disk-shaped scaffolds following in vivo implantation. Paraffin embedded sections (5 μm thick) were stained with eosin-hematoxylin (H&E) and Masson’s Trichrome and bright-field images of stained sections were acquired using an Axiovert 200M microscope (Carl Zeiss MicroImaging Inc.) with a 5X (0.16 NA) and 20X (0.8 NA) objective to evaluate construct morphology, remodeling and cell infiltration. To further assess cellular penetration into the implanted collagen disks, paraffin embedded sections were deparaffinized by washing the slides four times (2 min each wash) in histology grade xylenes (534056, Sigma-Aldrich) followed by serial rehydration with decreasing concentration of ethanol (100%, 95%, 70% and 0%). Rehydration steps were performed for 1 min each. The slides were then stained with DAPI (IC15757401, VWR) at a 1:200 dilution to visualize cell nuclei and imaged using large tile scans through the full thickness of the samples on a Zeiss

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700 LSM confocal microscope with a 20X (0.80 NA) objective. A custom ImageJ macro was used to analyze the nuclei distribution throughout each sample. Images were cropped to only contain the implanted collagen construct and then segmented using a binary mask to visualize the nuclei. Nuclei were identified using the built-in particle tracking tool and assigned a depth from the skin facing side of the construct. These nuclei counts were then assigned to binned depths to generate a histogram of nuclei count versus penetration depth. Large tile scan images of the lectin-labeled vascular network were acquired on a Zeiss 700 LSM confocal fluorescence microscope with a 20X (0.80 NA) objective. Images evaluating vessel depth and the presence of red blood cells were acquired on a Leica SP5 multiphoton imaging system with a 25X (0.95 NA) objective using 830 nm excitation with the InSight X3 laser and a 491 nm dichroic mirror for collection with Leica HyD detector. Contractile Human Ventricle – Fabrication and Characterization A model of the left ventricle corresponding in size to an early-stage fetal heart was designed as an open ellipsoidal shell with a 6.6 mm maximum outer diameter and 8 mm length from base to apex. The wall consisted of an inner 300 µm thick collagen layer, a middle 450 µm thick layer of cardiomyocytes and an outer 150 µm collagen layer (Fig. 3B). To print the ventricular construct, the multi-material 3D printing set-up was utilized (fig. S1E). Two syringes were used, one containing HES3 ESC-derived human cardiomyocytes, mixed with 0, 2, or 5% human ventricular cardiac fibroblasts (CC-2904, Lonza) and 20 mg/mL fibrinogen as a cellular bioink, and a second with acidified collagen bioink at 24 mg/mL. Syringes were loaded into Replistruder 3 extruders and mounted to the printer in a biosafety cabinet. Needle tips for both syringes were then aligned for dual-material printing to ensure high fidelity printing. The FRESH v2.0 support material was produced as previously described, using CDM3 media containing 10% FBS, 2 μM thiazovivin (S1459, Selleck Chemicals) and thrombin at 1 U/mL during the final wash prior to the final centrifugation step. Thrombin was used to polymerize soluble fibrinogen within the cell bioink to create a fibrin gel, which in combination with the collagen shells of the ventricle spatially maintained position of the printed cell bioink. The thrombin-containing FRESH v2.0 support bath was centrifuged to form compacted slurry and loaded into a 24 well tissue culture plate. The ventricle was printed using concentric perimeters at a speed of 3 mm/s for the first 1.5 mm height of the print, starting at the apex, and 10 mm/s for the remainder of the print. Printed ventricles were transferred to a 37ºC incubator to release the prints, and after 1 hour the melted gelatin support bath was exchanged with CDM3 culture media containing 10% FBS. Printed ventricles were cultured with media exchanges every two days, for up to 28 days. Electrophysiology of the printed ventricles was analyzed using calcium imaging during spontaneous and stimulated contractions. Printed human ventricles cultured for 14 days were washed with Tyrode’s solution to remove excess media and incubated in Tyrode’s solution containing 5 µM calcium indicator Cal 520 AM (21130, AAT Bioquest) and 0.025% Pluronic F127 (P2443, Sigma Aldrich) for 60 min at 37ºC, followed by 30 min at room temperature. Ventricles were then washed and incubated for 30 min in Tyrode’s solution to remove unbound dye. For imaging, ventricles were maintained in Tyrode’s solution at 37ºC using a custom heated stage. High-speed imaging of calcium transients up to 100 frames-per-second was performed using an epifluorescent stereomicroscope (SMZ1000, Nikon), GFP filter, X-Cite lamp (Excelitas), and Prime 95B Scientific CMOS camera (Photometrics). Calcium transients from spontaneous contractions were post-processed with custom MATLAB code that utilizes calcium signal peak

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times and cross-correlation to determine propagation delays across the tissue and calculate conduction velocity. Paced contractions were performed using two parallel platinum electrodes immersed in the Tyrode’s medium surrounding the tissues. An 80V, 20 ms, square wave pulse at 1, 2, and 4 Hz were applied with a Grass Stimulator. Point stimulation of ventricles were performed using a concentric biopolar microelectrode (30202, FHC Inc.) composed of an inner platinum-iridium pole an outer stainless-steel pole. The microelectrode was placed in contact with the ventricle surface and a 20V, 20 ms square wave pulse at 1 and 2 Hz was applied with a Grass Stimulator. Contractility of the printed ventricles was measured based on dimensional changes of the ventricular wall. To measure wall thickening of the printed ventricle during contractions, the inner and outer perimeters of the printed ventricle were manually outlined during peak diastole and systole across several contraction cycles acquired from a top-down view of the Cal520 stained ventricle. The Feret diameter of both the inner and outer walls was measured in ImageJ, and a wall thickness was calculated as half the difference of the outer and inner diameter. To track motion of the ventricular wall during contraction, the Imaris spot tracking algorithm was used to measure displacement of the inner, middle and outer wall regions and calculate displacement vectors. Percent change in area of the ventricle opening was calculated from top-down images of the ventricles by normalizing ventricle opening area during systole to ventricle opening area during diastole. Following functional imaging, printed ventricles were fixed and immunofluorescently stained. Ventricles were fixed in 4% formaldehyde (15710, Electron Microscopy Sciences) in PBS with 1:200 Triton X-100 (Thermo Fisher Scientific) for 1 hour followed by 3 washes in 1X PBS for 30 min. The ventricles were then blocked in 1% bovine serum albumin in PBS overnight and washed in PBS 3 times for 30 min each. Ventricles were stained with mouse anti-sarcomeric α-actinin primary antibody (A7811, Sigma-Aldrich) and rabbit anti-connexin 43 primary antibody (C6219, Sigma Aldrich) overnight at 1:100 dilutions at room temperature on a rotary shaker, then washed 3 times for 30 min in PBS. Then goat anti-mouse secondary antibody conjugated to Alexa-Fluor 488 (A28175, Life Technologies) and goat anti-rabbit secondary antibody conjugated to Alexa-Fluor 555 (A21428, Life Technologies) at a 1:100 dilution, DAPI (IC15757401, VWR) at a 1:200 dilution and phalloidin conjugated to Alexa-Fluor 633 (A22284, Life Technologies) at a 3:200 dilution were used to stain the ventricles overnight at room temperature, and then washed 3 times in 1X PBS prior to imaging. Tile scans and 3D z-stacks of the ventricles were acquired using a 20X (0.80 NA) and 63x (1.40 NA) objective on a Zeiss 700 LSM laser scanning confocal microscope and using a 25x (1.1 NA) objective on a Nikon A1R MP+ mutltiphoton confocal microscope.

Collagen Trileaflet Heart Valve – Fabrication, Functional Testing and Characterization Initial studies were performed on FRESH v2.0 printed compression test cylinders to identify fixation methods and infill percentages that produced collagen constructs with mechanical properties appropriate for use in a functional trileaflet valve. Mechanical characterization of printed alginate and collagen was performed using unconfined uniaxial compression testing. Compression test cylinders 10 mm in diameter and 5 mm in height were printed from either 23 mg/ml acidified collagen or 4% (w/v) alginic acid. A support bath of 0.10% (w/v) CaCl2, 25 mM HEPES, pH 7.4 was used to enable printing either bioink into the same support bath using a 150

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μm needle. Compression cylinders (N = 6 of each type) were printed at 35% (low), 50% (medium), or a near-solid infill of 75% or 90% (high) for alginate and collagen, respectively, using a 60 μm layer height. Following printing, constructs were heated to 37ºC to melt the gelatin, and placed in a rotary incubator (Innova 42R, New Brunswick) at 40oC, 60 RPM in 1X PBS for 24 hr to ensure no interstitial gelatin remained trapped within the constructs. After complete removal of interstitial gelatin, compression cylinders were placed in 50 mL fixative solutions. All fixing solutions (except 1X PBS) were buffered to pH 7.4 with 25 mM HEPES. Alginate samples were placed in 1% (w/v) CaCl2. Collagen samples were fixed in 1X PBS (control), 75% (v/v) ethanol with 0.05% (v/v) glutaraldehyde (GA), or 75% (v/v) ethanol with 0.5% GA under static conditions at room temperature for 7 days. The diameter of each cylinder was measured before mechanical testing. Compression testing was performed on an Instron 5943 at a strain rate of 1 mm/min until approximately 60% strain. The elastic modulus of each sample was calculated from the slope of the linear elastic region of the stress-strain curves. The printed trileaflet heart valve was based on a digital model downloaded from the NIH 3D Print Exchange website (https://3dprint.nih.gov/discover/3dpx-000452). Excess regions of the model proximal and distal to the center of the valve were removed using Meshmixer and scaled to final dimensions of 27.8 mm diameter and 19.96 mm height. An isolated leaflet model was generated in Meshmixer by removing the valve wall and the other two leaflets. G-code was generated using Slic3r with print settings of 60 μm layer height, 2 perimeters, 50% infill, 23 mm/s movement speed, and a 150 µm ID stainless steel needle. Prior to printing, the gelatin support bath was processed as previously described with a washing solution of 0.10% (w/v) CaCl2, 25 mM HEPES, pH 7.4 and transferred into a print container ~50% larger than the valve. Alginate and collagen were printed using the same process as for the compression test cylinders and released and washed in the same manner to remove gelatin. For collagen valves, compression testing established that constructs printed at 50% infill and fixed with 0.5% GA achieved compressive modulus comparable to alginate constructs (fig. S13A), leading to all collagen valves to be printed and subsequently fixed with these conditions. Printed collagen valves underwent an intermediate fixation step by placing them into 1X PBS, 0.5% GA for 24 hr at room temperature after the overnight rotary incubation step. Afterwards, valves were then transferred into 75% (v/v) ethanol with 0.5% GA for 6 days to prevent infection and ensure thorough fixation. A water column was used to assess the maximum transvalvular pressure that the 3D printed valves could support prior to failure. The water column consisted of a custom designed valve adapter inserted into a 1.25” ID, 4’ long clear plastic tube (McMaster-Carr). The valve adapter was designed in SOLIDWORKS to precisely fit the printed trileaflet valve and 3D printed from NinjaFlex (NinjaTek) flexible filament to hold the valve while forming a water-tight seal at the bottom of the water column. To measure the maximum transvalvular pressure, the printed valves were placed in the water column at the base and water was steadily added above the valve until the point of valve failure, which occurred when water began to pass through the valve. The maximum height of the water before failure was recorded and used to calculate the maximum pressure. The function of the collagen valves was assessed using a benchtop system that mimicked physiological pressure and pulsatile flow. Experiments were performed on a flow loop using a pulsatile pump (Pulsatile / Blood Pump 1421, Harvard Apparatus), where the mechanical ball

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valve on the pump outlet was replaced with a printed collagen valve. A two-part housing for the collagen valve was 3D printed from polycarbonate (PC-Max, Polymaker), which bolted together around the valve and had integrated hose barbs to connect to the pulsatile flow loop. An ultrasonic Doppler flow sensor (SONOFLOW CO.55/080 V2.0, SONOTEC) was placed at the entrance to the valve housing to assess valvular regurgitation while a Luer lock pressure transducer recorded pressure data (EMKA iox2, EMKA Technologies). A penrose drain (DYND50426, Medline) was placed distal to the valve to provide compliance to the system and an endoscope with integrated digital camera (Depstech) was inserted distal to the valve housing via a T junction in the flow loop to record opening and closing of the valve within the housing. A solution of 40% (v/v) glycerol (Thermofisher Scientific) was used as a blood analogue. To mimic the lower transvalvular pressures in the right side of the heart, pump settings were set to a stroke volume of 30CC, 30 BPM with a systole time of 33%, resulting in pressures of approximately 25/15 mmHg when using mechanical ball valves. The mechanical outlet valve was then replaced with a collagen valve and pumped until failure using the same settings. Valve failure was deemed to be 40% regurgitation. The area between the leaflets during maximum opening was quantified in ImageJ and compared to the inner diameter of the valve at the base of the leaflets to calculate percent opening. For measuring valve print fidelity, 3D gauging of collagen valves was performed. Printed collagen valves with barium sulfate added for X-ray contrast were glutaraldehyde-fixed, embedded in agarose, and imaged with μCT. Scans were performed on a VivaCT 50 (SCANCO Medical AG) with 10.3 µm voxel resolution, 70 KVp beam energy, 85 mA current intensity, and 100 ms exposure. The 3D reconstruction was performed using SCANCO software from the raw files and the resulting volumes were processed with the SCANCO microCT 3D morphometry software. DICOM data from the μCT scan was processed in ImageJ to remove excess noise (despeckle). The data was then converted to a surface STL file using InVesalius (CTI). The scan STL was then overlaid with the original in Geomagic Wrap (3D Systems). Quantification of surface deviation between the scan and original was performed using the Deviation Analysis tool to generate a heatmap of overprinted and under-printed regions. To assess cellular compatibility of collagen valve constructs, collagen leaflets were printed, released, and then left unfixed or fixed as previously described for the full valve. Leaflets were washed overnight with 1X PBS on a rotary shaker at 60 RPM, followed by 2 more 1X PBS washes of 1 hr to remove any remaining GA and ethanol (if present). Leaflets were then treated with UV-Ozone for 15 min to sterilize prior to seeding. HUVECs were seeded at 2 X 105 cells/cm2 onto GA-fixed (75% (v/v) ethanol, 0.5% (v/v) GA, 25 mM HEPES, pH 7.4) and unfixed (75% (v/v) ethanol, 25 mM HEPES, pH 7.4) Printed collagen heart valve leaflets as well as fibronectin-coated (50 μg/mL) glass coverslips were used to assess monolayer formation (N = 3 of each type) with media changed every 24 hr. At day 2 and 5 HUVEC seeded coverslips and valve leaflets were washed with 1X PBS with Ca2+ Mg2+ three times and fixed for 15 min with 4% (v/v) formaldehyde. After three washes with 1X PBS, HUVECs were blocked and permeabilized for 1 hr at 37ºC in PBS with 0.05% (v/v) Triton X-100, 50 mM glycine, and 5% (v/v) FBS. Primary antibodies were added for DAPI (1:200), Alexa Flour 555 conjugated Phalloidin (1:500, Life Technologies), and mouse anti-CD31 (1:500, 3528S, Cell Signal) and incubated overnight at 4ºC. Following primary antibody incubation, samples were washed three times with 1X PBS and stained with Alexa Fluor 488 conjugated goat anti-mouse secondary antibody (1:500, Life Technologies) for 1 hour.

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Confocal fluorescence imaging was performed using a 20X objective (0.80 NA) on a Zeiss 700 LSM microscope.

MRI-Templated Multiscale Vasculature – Computational Generation, Fabrication, Perfusion and Characterization A subsection of the 3D heart model containing the left ventricle and coronary vessels was extracted to serve as a template to generate a multiscale vessel network in a collagen scaffold derived from MRI data (fig. S14A). The left ventricle and corresponding coronary arteries from the full heart 3D model were each imported into Meshmixer and were dilated in order to eliminate gaps between the ventricle surface and arteries when combined. The combined mesh was further dilated and saved as an OBJ file for later use as a masking template, while the original mesh was exported in OBJ format and imported into Element software (nTopology) to be used for vasculature generation. Element software was used to generate multiscale vasculature within a segmented volume, corresponding to the wall of the left ventricle, by creating a stochastic Voronoi distribution of nodes and beams. The node density and beam diameter were varied based on absolute orthogonal distance from another volume, the left coronary arteries. The coronary arteries model was imported and a 3D heatmap (fig. S14B) indicating distance from the arteries was generated. This 3D distance heatmap guided the generation of two modifier surfaces, one modifying node density and another beam diameter. For node density modification, regions immediate to the arteries (red) produced a less dense lattice while those further away (blue) produced a locally denser lattice, however all lattice beams were the same diameter throughout. The second beam diameter modifier was then used to adjust the diameter of the lattice beams based on distance from the arteries with the beams closer (red) to the arteries being thicker and beams further away (blue) from the arteries being thinner (fig. S14, C and D). This two-step lattice generation of density followed by diameter produced a 3D space-filling vasculature within the ventricular wall. This vasculature model was exported as an OBJ, imported into Meshmixer, and Boolean-added to the coronary arteries file to produce a complete multiscale vasculature. Finally, this complete multiscale vasculature was subtracted from the previously created positively-offset combined mesh to produce the final vascularized left ventricle model with hollow internal channels. A sub-region of the vascularized left ventricle model containing the left anterior descending (LAD) artery and internal microvasculature was then selected and exported as an STL file (fig. S14, E to J). The model was manually repaired or thickened in thin spots to avoid printing errors such as overly thin walls and arterial channels with holes. The model was FRESH v2.0 printed using 24 mg/mL acidified collagen bioink and fixed using settings and methods identical to the neonatal-scale collagen heart (fig. S14K). Perfusion of the printed LAD sub-region was performed by flowing 40% (w/v) glycerol (Thermo Fisher Scientific) as a blood substitute colored with red food coloring into the main branch of the LAD with an 840 µm ID needle and a syringe pump. A custom construct mount was 3D printed from PLA to hold the construct off the container bottom during the perfusion experiments to allow fluid to drain properly. Video of the perfusion of the multiscale vasculature was acquired on a Canon 70D DSLR camera. To visualize the internal vasculature, including vessels down to the 100 µm length scale, the construct was cleared via serial ethanol dehydration followed by submersion into a 2:1 mixture of benzyl benzoate (Sigma-Aldrich) and benzyl alcohol (Sigma-Aldrich)

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(BABB) for 24 hr (fig. S14, L and M). The construct was again perfused with the glycerol blood substitute and the perfusion was recorded (fig. S14, N to P). Neonatal-Scale Collagen Heart – Fabrication and Characterization The 3D MRI-based digital model for the human heart was downloaded from the BodyParts3D database (http://lifesciencedb.jp/bp3d/) (24), repaired to remove mesh defects, and re-sized to neonatal scale using Meshmixer. The model was then exported as an STL and sliced into G-code using Slic3r based on a 150 µm ID needle and 80 µm thick layers. The acidified collagen bioink at 24 mg/mL was loaded into a 10 mL Hamilton syringe and mounted onto a Replistruder 3 syringe pump extruder. The FRESH v2.0 support bath was processed using a washing solution of 50 mM HEPES at pH 7.4 for the final centrifugation and transferred into a print container approximately 50% larger than the construct. To increase X-ray radiocontrast for µCT imaging, some prints were performed using collagen bioink supplemented with 2.0% (w/v) barium sulfate. After printing, collagen hearts were released from the FRESH support bath by melting the gelatin at 37ºC. Due to the large size of the heart, optical imaging was performed using a digital SLR camera (Canon 70D, Canon) under oblique illumination. For µCT imaging, collagen hearts with barium sulfate were fixed in 1% (v/v) formaldehyde, 25 mM HEPES, pH 7.4 for 24 hr and embedded in 4% (w/v) agarose inside the holder provided by the manufacturer. Scans were performed on a VivaCT 50 (SCANCO Medical AG) with 10.3 µm voxel resolution, 70 KVp beam energy, 85 mA current intensity, and 100 ms exposure. The 3D reconstruction was performed using SCANCO software from the raw files and the resulting volumes were processed with the SCANCO microCT 3D morphometry software. For additional image processing and 3D rendering, image slices were background subtracted and cropped within ImageJ, and 3D surface rendering and movie creation was performed using Imaris. STATISTICAL ANALYSIS Statistical analysis was performed using Prism 6 (GraphPad Software) software, and appropriate statistical tests were chosen according to experimental conditions and data requirements. Each experiment was performed in triplicate unless otherwise specified. For comparison of particle Feret diameter between FRESH v1.0 and FRESH v2.0 (Fig. 1F) and maximum transvalvular pressure (Fig. 4J), Student’s t-tests were performed. For statistical analysis of the percent wall thickening in contracting ventricles (Fig. 3M) and compressive modulus of FRESH printed constructs (fig. S14A), one-way analysis of variance (ANOVA) followed by Tukey’s multiple comparisons post-test was performed. Statistical significance was considered a p-value of <0.05. For statistical analysis of nuclei penetration depth of implanted constructs (Fig. 2L and fig. S8B), two-way ANOVA was performed with independent variables fabrication method and nuclei depth and a dependent variable nuclei count. Bonferonni multiple comparisons post-test was performed to determine statistical differences of nuclei count between cast and printed samples at each depth. For statistical analysis of viability in perfused versus non-perfused constructs (Fig. 2F), a two–way ANOVA was performed with independent variables of perfusion status (i.e. perfused or non-perfused) and depth and a dependent variable percent viability. Bonferonni multiple comparisons post-test was used to determine statistical differences between perfused and non-perfused samples at each depth.

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Fig. S1. Modification of a thermoplastic 3D printer for FRESH v2.0 3D bioprinting using an open-source syringe-based extruder. (A) Rendering of the assembled Replistruder 3 syringe-pump extruder. (B) The Replistruder 3 mounted onto a PrintrBot Simple Metal 3D printer. (C) Rendering of three assembled Replistruder 3’s mounted on the multi-extruder carriage, allowing the use of up to 3 separate bioinks for multi-material printing. (D) The multi-extruder carriage and Replistruder 3’s mounted on a FlashForge Creator Pro 3D printer. (E) The multi-extruder system with two Replistruder 3’s mounted in the carriage. The needle tip of each extruder is aligned to the laser-cut alignment guides in the base plate using the position adjustment screws located around the syringe body and the vertical adjustment screws (not shown). Printing is performed in the center Petri dish containing the FRESH support bath with the outer Petri dishes holding DI water to keep the inactive needles clean and to prevent clogging. (F) Two Replistruder 3 extruders mounted on a MakerBot 3D printer and placed inside a biosafety hood for sterile FRESH printing.

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Fig. S2. Using centrifugation speed to control FRESH v2.0 support bath rheology and printed collagen filament morphology. (A) The storage modulus G’ as a function of strain for support baths centrifuged at varying speeds. The plateau of G’ and the critical strain where the support bath yields from a solid to liquid state both increase with increasing centrifugation speeds. (B) Phase contrast and confocal reflectance images of collagen filaments FRESH printed into support baths centrifuged at various speeds. Note that a centrifuge speed of 2,000 g was found to be the optimal setting for microscale filament morphology and filament-to-filament adhesion required for 3D printing larger constructs.

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Fig. S3. Role of FRESH v2.0 support bath fluid-phase buffer composition on collagen filament morphology. (A) Collagen filaments were FRESH printed with an 80 µm inner diameter needle into support baths with different buffering solutions and released immediately after printing (top row, 15 min) or released 24 hours after printing (bottom row). The presence of small fibrous features surrounding the primary printed filaments in several 24 hour conditions (1x PBS, 50 mM HEPES and 9 mM CaCl2) indicate incomplete gelation of extruded collagen when initially printed. (B) 3D surface rendering of confocal reflectance imaging of collagen filaments printed into support baths with different buffering solutions and released after 24 hours. Printed filaments show a variety of surface morphologies, from fibrous (1x PBS, Essential 8, extracellular buffer), to smooth (10x PBS), to indented (25 and 50 mM HEPES), to a combination of both (50 mM HEPES and 1.37 M NaCl at pH 7.4, and 50 mM HEPES and 1.37 M NaCl at pH 8.2). These surface morphologies result from differences in gelation kinetics and rheology of the support bath. The support bath used for studies in this manuscript consisted of pH 7.4, 50 mM HEPES.

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Fig. S4. 3D Bioprinting four unique bioinks with distinct cross-linking mechanisms within the same FRESH 2.0 support bath. (A to D) Collagen, alginate, fibrinogen and MeHA inks individually printed into separate containers filled with a FRESH v2.0 support bath with HEPES, CaCl2, and thrombin demonstrates the ability of one support bath to enable printing of biopolymers that crosslink and gel due to pH, ionic, enzymatic and light-sensitive mechanisms. (E) A rectangular frame with the letters “CMU” cut out printed from collagen (red). The letters “C”, “M”, and “U” were printed from alginate (blue), fibrinogen (magenta) and MeHA (green), respectively, into the cutouts in the frame.

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Fig. S5. Perfusion setup for collagen tubes. (A) A CAD model of the perfusion chamber. (B) The perfusion chamber 3D printed using stereolithography. (C) The 3D printed perfusion chamber (clear yellow plastic) with the 3D printed polylactic acid (PLA) negative mold (gray plastic) placed on top. Polydimethylsiloxane (PDMS) is poured in between these pieces and cured to form a well to hold the collagen tube construct. (D) The complete perfusion chamber setup with a central PDMS well and perfusion needles inserted. (E) The FRESH printed collagen tube construct within the perfusion chamber showing insertion of perfusion needles through both ends. (F) The perfusion chamber with media added and lid placed for sterile culture. (G) The perfusion chamber circuit used for long term culture including a perfusion chamber with collagen tube construct, multichannel peristaltic pump and media reservoir. The entire setup is then placed within a cell culture incubator.

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Fig. S6. Dimensions and quantification of perfusion of fluorescent dextran through FRESH v2.0 printed collagen tubes. (A) 3D schematic of a 1.4 mm diameter collagen tube within a collagen frame for in vitro perfusion studies. (B) Cross-section of the collagen tube imaged using confocal reflectance. (C) The collagen tube perfused with red dye to show patency. (D) Kymograph of the fluorescent intensity of 10 kDa fluorescent dextran perfused within the printed collagen tube as shown in (C) and diffusing through the wall into the surrounding space. (E) Relative diffusion of 10 kDa and 70 kDa dextran through the tube wall from the kymograph (D, white box).

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Fig. S7. In vivo implantation of casted and FRESH v2.0 printed collagen disk constructs. (A) Renderings and images of casted and FRESH v2.0 printed collagen constructs for implantation. (B) Collagen constructs were subcutaneously implanted on the dorsal-side of C57BL6 immuno-competent mice. (C) Timeline of construct implantation and harvest at 3, 7 and 14 day timepoints. (D) Whole collagen constructs after harvesting at respective time-points. (E) Histology at 3, 7 and 14 date timepoints stained by H&E and Masson’s Trichrome. Dashed line indicates construct interface with surrounding tissue. Note the increased cell infiltration in FRESH v2.0 printed constructs compared to casted controls.

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Fig. S8. Analysis of cell penetration into casted and FRESH v2.0 printed collagen disk constructs following in vivo implantation. (A) Collagen disk constructs casted or FRESH v2.0 printed were implanted in vivo subcutaneously in mice and then explanted at 3 and 7 days. Representative fluorescence images show cellular penetration of DAPI-stained nuclei (white) through casted and printed disks. (B) Mean cell density as a function of penetration depth into the cast and printed collagen disks following 3 days of implantation, showing greater cell infiltration through the entire thickness of the printed construct (N = 6).

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Fig. S9. Vascularization after 10 days in vivo implantation of casted and FRESH v2.0 printed collagen disk constructs functionalized with fibronectin and VEGF. (A) Histology of casted and FRESH v2.0 printed collagen constructs without VEGF explanted after 10 days, stained with H&E. Dotted lines indicate construct border. (B) Histology of casted and printed collagen constructs with VEGF explanted after 10 days, stained with H&E. Dotted lines indicate construct border.

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Fig. S10. Vascularization of a FRESH v2.0 printed collagen disk construct functionalized with fibronectin and VEGF. (A) A large area confocal tile scan of a FRESH v2.0 printed collagen disk showing vascularization from the host into the construct, stained for nuclei (DAPI, blue) and perfused blood vessels (fluorescently labeled lectin via tail vein injection, red). (B) Zoomed in image of the small capillary-scale vessel outgrowth at the edge of the vascular network into the surrounding collagen. (C) Zoomed in image of a region with varying vessel size and density within the vascular network.

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Fig. S11. Viability of cells and structure of FRESH v2.0 printed ventricle. (A) Dual material FRESH v2.0 printed sheet with distinct high-fidelity regions of cells and collagen. (B) LIVE/DEAD staining from a subregion in (A) shows high cell viability within the cell ink 1-hour post-printing. Ventricles printed with a cardiomyocyte bioink containing (C) 0%, (D) 2%, and (E) 5% fibroblasts show significant compaction of the 5% fibroblasts ventricles and minimal structural deformation from cell-driven compaction in the 0% and 2% fibroblast ventricles. Calcium staining and imaging of ventricles reveal (F) patchy cardiomyocyte distribution in 0% fibroblast ventricles with contractility, (G) homogenous coverage of cardiomyocytes in 2% fibroblast ventricles with high contractility, and (H) homogenous coverage of cardiomyocytes in 5% fibroblast ventricles with minimal contractility. (I) Immunofluorescent staining of the ventricle wall from a laterally sectioned 2% fibroblast ventricle fixed at day 14 shows intact printed collagen sections (magenta, second harmonic generation) surrounding a dense layer of cardiomyocytes (α-actinin, green; nuclei, blue). High magnification image of the printed cell-section from (J) single and (K) tiled field of views reveal the high interconnectivity of the dense layer of cardiomyocytes.

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Fig. S12. Spontaneous conduction patterns of FRESH v2.0 printed ventricles. (A) FRESH v2.0 printed ventricle displayed stable, spontaneous pinned rotor following a period of field stimulation, and (B) calcium mapping of the subregion (A, yellow box) confirming rotor activity. (C) Multiple calcium waves were observed propagating in a FRESH printed ventricle prior to stimulation, and (D) calcium mapping from a top-down view of the ventricle showing distinct separation of two propagating waves around the walls of the ventricle.

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Fig. S13. Heart valve mechanical analysis, surface gauging, and HUVEC monolayer culture. (A) Compressive elastic modulus of FRESH v2.0 printed alginate and collagen disks printed at low (35%), medium (50%) and high (75% or 90%) infill percentages. Collagen was either unfixed or fixed with 0.05% or 0.5% glutaraldehyde. Statistically significant differences were shown between most conditions. The graph indicates a not significant (n.s.) difference in the elastic modulus between medium infill alginate and medium infill collagen with 0.5% (v/v) glutaraldehyde fixation, the two conditions that produced functional valves in our studies (N = 6, data are mean ± SD, *** indicates P < 0.001, **** indicates P < 0.0001 by one-way ANOVA followed by Tukey’s post-hoc test. (B) Gauging of the printed collagen heart valve imaged in 3D using μCT and compared to the original 3D STL files. Direct overlay of the μCT data (grey) against the original STL file (blue). (C and D) The differences between the μCT data relative to the STL are shown as overprinting (yellow to red) and under-printing (light to dark blue) from (C) top-down and (D) side views. (E) Human umbilical vein endothelial cells (HUVECs) seeded onto glass coverslip controls and unfixed printed collagen valve leaflets and cultured for 2 and 5 days, stained for nuclei (blue), F-actin (red), and CD-31 (green), demonstrating confluent monolayer formation.

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Fig. S14. Computational generation, 3D printing and perfusion of MRI-templated multiscale vasculature. (A) A 3D model of the heart from patient-specific MRI dataset with the coronary arteries (red) and veins (blue) of the left ventricle (purple). (B) The heatmap on the surface of the left ventricle shows the distance from the coronary arteries, scaling from red to blue. For this heat map, as the distance from the coronary arteries increases the size of the generated vessels will decrease while the volume density of vessels will increase. (C) Creation of the microvascular network using the generated density map from (B) and a Voronoi space-filling algorithm, extended throughout the ventricular wall. (D) A magnified view of the generated 3D microvasculature anastomosed to the left coronary artery. (E) A solid sub-region (brown) of the ventricular wall selected for 3D printing that includes the computationally generated vascular network from the left coronary artery. (F) A zoomed in view showing the left anterior descending (LAD) coronary artery sub-region. (G) An x-ray view of the LAD sub-region where the vessel network is rendered as void space. (H) Isolation and pseudo-colored perfusion (red) of the LAD artery through the sub-region. (I) Pseudo-colored perfusion (red) of the LAD artery and generated microvasculature. (J) A 3D render of the LAD sub-region for 3D bioprinting using FRESH v2.0. (K) The LAD sub-region FRESH printed using collagen. (L) The LAD sub-region shown midway through the tissue clearing process using benzyl alcohol benzyl benzoate (BABB) to improve visualization of internal vessels during perfusion studies. (M) Fluorescence image of the printed LAD sub-region showing the printed collagen structure and internal vessel paths. (N-P) Perfusion of glycerol-based blood substitute (red) through the BABB cleared FRESH printed construct showing (N) the cleared non-perfused sample, (O) perfusion through large diameter vessels, and (P) perfusion through the final small diameter vessels respectively. Note that vessels at the periphery of the sub-region were open to the surrounding water bath, accounting for the diffusive spreading of the blood substitute.

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Fig. S15. Organ-scale imaging of the FRESH v2.0 printed collagen heart. (A) Top-down photograph of the FRESH v2.0 printed collagen heart showing aortic valve and pulmonary artery. (B) Posterior photograph showing the right pulmonary veins and inferior vena cava. (C) Left-lateral photograph showing the left pulmonary veins and branches of the left coronary artery. (D) Posterior view of the printed collagen heart rendered from µCT data in Imaris showing the right pulmonary veins and the 3D printed infill pattern. (E) Cross-sectional view of the left and right ventricles from the µCT data revealing the internal trabeculae and 3D printed infill pattern within the ventricular walls. (F) Cut away view of the heart highlighting the aortic valve and the aortic root. (G) Top-down cut away view showing the aortic and pulmonary valves, left and right atrium, and left atrial appendage. (H) Coronal cross-section showing both ventricles and the aortic valve. (I) Cross-sectional view showing the right atrial appendage, left and right ventricles, pulmonary valve, and internal trabeculae within the left ventricle.

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Table S1. Media formulations used in culture of human embryonic stem cells and differentiation into cardiomyocytes. MEDIA COMPONENTS RPMI/B27 • RPMI-1640 medium (21870076, Thermofisher)

• 1% v/v L-glutamine (25030081, Thermofisher) • B27 supplement (17504044, Thermofisher)

CDM3 • RPMI-1640 medium (21870076, Thermofisher) • 1% v/v L-glutamine (25030081, Thermofisher) • 500 µg/mL human albumin (A9731, Sigma) • 213 µg/mL L-Ascorbic acid 2-phosphate sesquimagnesium salt hydrate

>95% (A8960, Sigma) CDM3L • RPMI-1640 lacking glucose (11879020, Thermofisher)

• 500 µg/mL human albumin (A9731, Sigma) • 213 µg/mL L-ascorbic acid-2-phosphate (A8960, Sigma) • 7.1 mM sodium-lactate (L4263, Sigma)

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Captions for Movies S1 to S10 Movie S1. FRESH v2.0 Printing CMU. Time-lapse sequence of 3D bioprinting the letters “CMU” from an Alcian blue-stained 4% (w/v) alginate ink using FRESH v2.0. Printing is performed at room temperature and raised to > 37oC after completion, melting the gelatin support bath and releasing the print. Alginate was used instead of collagen as it provided higher contrast within the support bath, making it easier to visualize the printing method. Movie S2. Perfusion of 10 kDa fluorescent dextran through the FRESH v2.0 printed collagen perfusion tube. Video shows the peristaltic flow of a 10 kDa fluorescein dextran through a 1.4 mm inner diameter collagen tube during the initial 2 minutes, sped up 6x. Fluorescence intensity has been pseudo-colored with the fire LUT. Movie S3. Flythrough of a FRESH v2.0 printed collagen construct with 1 mm diameter printed pores and ~30 µm diameter micropores due to removal of the gelatin microparticle support bath. Confocal reflectance microscopy of a printed collagen construct reveals an extensive and uniform porous network purposely generated by the incorporation of the gelatin microparticles into the construct during the printing process and removal during the release process. Movie S4. 3D rendering of vascularization in the FRESH v2.0 printed VEGF doped collagen construct after 10 days in vivo. 3D visualization and flythrough of a printed collagen construct showing a multiscale vascularized network (lectin, red) and nuclei (DAPI, blue) down to vessels as small as 5 µm in diameter. Movie S5. In vivo vascularized network within the FRESH v2.0 printed collagen construct contains red blood cells. Multiphoton imaging at 870 nm excitation of a printed collagen construct following implantation into mice. Vessels are seen at depths of up to 200 µm and contain red blood cells. Movie S6. Calcium imaging of FRESH v2.0 printed human ventricle during spontaneous and paced contractions. Ventricles were FRESH v2.0 printed using human stem cell-derived cardiomyocytes and adult human cardiac fibroblasts. Shown in sequence are (i) side views of a ventricle during spontaneous and 1 Hz and 2 Hz field stimulated contractions. (ii) Top-down view of a ventricle during spontaneous and 1 Hz and 2 Hz field stimulated contractions. (iii) Side view of blebbistatin treated ventricle during 1 Hz point stimulation with a bipolar electrode followed by an overlaid time delay map of conduction. (iv) Top-down view of ventricle during spontaneous contractions showing two simultaneous wave fronts followed by an overlaid time delay map of conduction. (v) Side view of ventricle during spontaneous contractions demonstrating pinned rotor-like electrical activity followed by an overlaid time delay map of conduction. Movie S7. Handling and pulsatile flow testing of FRESH v2.0 printed collagen trileaflet heart valve. A trileaflet valve 28 mm in diameter is handled after fixation with glutaraldehyde. The valve is capable of sustaining its weight in air and is printed with high enough fidelity to prevent the fusion of the leaflets, allowing for individual actuation. The valve is then placed in an in vitro pulsatile flow loop and imaged using an endoscope. The pressure in the system was lowered to

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25/15 mmHg by adjusting the stroke rate and volume to 30 BPM and 30CC, respectively, to approximate the physiologic pressures experienced by the pulmonary valve in vivo. Movie S8. FRESH v2.0 printed collagen trileaflet heart valve imaged by μCT. A 3D reconstruction of the trileaflet heart valve printed from collagen with 2% (w/v) barium sulfate. Incorporating barium sulfate into the ink increases X-ray contrast and allows for the internal geometry of the valve to be visualized. The entire valve structure is panned through first in the xz plane followed by top-down through the xy plane. Movie S9. Computational generation, 3D printing and perfusion of MRI-templated multiscale vasculature. Movie depicts the design and fabrication process of the generative vasculature followed by perfusion of a glycerol blood substitute through the printed vascular network, with vessels ranging from 5 mm down to ~100 µm in diameter. Movie S10. 3D rendering and digital sectioning of µCT of the FRESH v2.0 printed collagen heart. An adult human heart from MRI was scaled to down to neonatal size 37 mm in diameter, FRESH v2.0 3D printed using collagen and imaged by µCT. Imaris image analysis software was used to import and visualize the µCT data (grey). Surface renderings of the heart were used to enhance visualization of the external printed features (tan). Digital sectioning of the printed heart revealed the internal architecture and detail contained within the print including the valves and trabeculae.

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