{'fraction_non_alphanumeric': 0.03482187272796847, 'fraction_numerical': 0.01970688401637776, 'mean_word_length': 4.894000902119982, 'pattern_counts': {'":': 0, '<': 1, '<?xml version=': 0, '>': 0, 'https://': 0, 'lorem ipsum': 0, 'www.': 0, 'xml': 0}, 'pii_count': 0, 'substrings_counts': 0, 'word_list_counts': {'cursed_substrings.json': 0, 'profanity_word_list.json': 0, 'sexual_word_list.json': 0, 'zh_pornsignals.json': 0}}
Biodegradable scaffold with built-in vasculature for organ-on-a-chip engineering and direct surgical anastomosis

We report the fabrication of a scaffold (hereafter referred to as AngioChip) that supports the assembly of parenchymal cells on a mechanically tunable matrix surrounding a perfusable, branched, three-dimensional microchannel network coated with endothelial cells. The design of AngioChip decouples the material choices for the engineered vessel network and for cell seeding in the parenchyma, enabling extensive remodelling while maintaining an open-vessel lumen. The incorporation of nanopores and micro-holes in the vessel walls enhances permeability, and permits intercellular crosstalk and extravasation of monocytes and endothelial cells on biomolecular stimulation. We also show that vascularized hepatic tissues and cardiac tissues engineered by using AngioChips process clinically relevant drugs delivered through the vasculature, and that millimeter-thick cardiac tissues can be engineered in a scalable manner. Moreover, we demonstrate that AngioChip cardiac tissues implanted via direct surgical anastomosis to the femoral vessels of rat hindlimbs establish immediate blood perfusion.

temperature and imaged with μCT, where a threshold was applied to the acquired image to reveal the built--in network against the other structure of the scaffold.
AngioChip scaffold degradation. To prepare the nano--porous scaffold samples, scaffolds (3mm×10mm×200μm) with 60% porogen content were UV cross--linked for 4 min at intensity of 10mJ/cm 2 s and weighted. To prepare the pore--free scaffold samples, scaffolds (3mm×10mm×200μm) with 0% porogen content were UV cross--linked for 10 min at intensity of 10mJ/cm 2 s and weighted. The degradation rates of the scaffolds were determined by tracking the scaffold mass loss over time in PBS (pH 7.4, 37°C) or in 0.1 M NaOH. After each incubation period, the samples were first thoroughly washed with distilled water and then dried. The mass loss was calculated by comparing the initial mass with the mass measured at the specific time point.
AngioChip burst pressure. To determine the burst pressure of the AngioChip scaffolds without micro--holes, the inlet of the scaffold was connected to a nitrogen tank with a pressure gauge using Tygon tubing, while the outlet of the scaffold was sealed with epoxy glue. Pressure was increased gradually until a leak was observed from the scaffold network and the peak pressure was recorded as the burst pressure. If a leak from the junctions at the inlet or the outlet was observed first, the recording was discarded.
AngioChip scaffold and myocardium mechanical testing. The mechanical properties of the AngioChip scaffold were measured in PBS with a Myograph (Kent Scientific) in the circumferential (along the long edge) and longitudinal direction (along the short edge) of the AngioChip scaffolds. Strain--to--failure values were determined from the strain at the breaking point along the curves. The effective elasticity of the adult rat myocardium was measured from the adult hearts of Lewis rats, sacrificed according to a protocol approved by the University of Toronto Animal Care Committee. The adult rat myocardium was sliced into 7mm long strips with a width of 2--4mm and thickness of 2--4mm along the circumferential direction or longitudinal direction of the heart.
The rat myocardium was only stretched up to 20% strain because the stress--strain curve of the myocardium becomes nonlinear outside of the 20% strain region as shown in other studies 1 . In addition, normal heart physiological strain from active contraction is only within 15% 2 . The AngioChip scaffolds were stretched until break, a strain much higher than 20%, in order to determine the ultimate tensile strength. The effective elasticity of both the scaffold and the myocardium was derived from the slope of the stress--strain curve between 0% to 10%, consistent with previous studies 1 (Supplementary Figure 9, 10).
Anisotropic ratio was determined by dividing the effective elasticity in the circumferential direction with the effective elasticity in the longitudinal direction. The three different scaffold designs (Figure 2a--c) had different lattice structures but the same built--in network with a wall thickness of 50μm, an inner luminal dimension of 50μm by 200μm for the inlet, outlet and first order branch, and an inner luminal dimension of 50μm by 100μm for the second order branch. The scaffold lattice was made of 50μm struts. In design A, the struts were spaced 250μm apart in the long--edge direction, 100μm apart in the short--edge direction, and they were continuous in the z--axis. In design B, the struts were spaced 250μm apart in the long--edge direction, 100μm apart in the short--edge direction, and 50μm apart in the z--axis. In design C, the struts were spaced 550μm apart in the long--edge direction, 175μm apart in the short--edge direction, and 50μm apart in the z--axis. AngioChip permeability. The permeability of large molecules (TRITC--Dextran, 70 kDa, Sigma) from the built--in network of the micro--hole free AngioChip, AngioChip with 10μm micro--holes, or AngioChip that was both endothelialized and contained 10µm micro--holes was measured. TRITC--Dextran was perfused at 0.7μL/min through the network with inlet concentration of 10 μM for 20hr. After 20 hr, the solution in the middle wells was collected and the concentration of fluorescent molecules was correlated to a standard curve determined with a fluorescence plate--reader. The channel permeability of the entire AngioChip scaffold to large molecules was determined from the net rate of diffusion (total accumulated fluorescent molecules in the middle chamber/time) and the luminal surface area of the network, using Fick's Law with the assumption that the average concentration of fluorescent molecules inside the network is the same to that at the inlet. To illustrate diffusion and metabolism of biomolecules in the tissue space, 100μM carboxyfluorescein diacetate (CFDA, 557 Da) in PBS was perfused at 0.7μL/min through the scaffold network with 10μm micro--holes surrounded by cardiac cells (1 day after seeding) for 60min.

POMaC material biocompatibility in vivo:
Poly--L--lactic acid (PLLA, 85,000--160,000Da, Sigma) was first dissolved in cholorform. POMaC pre--polymer was prepared as described above. PLLA solution and POMaC solution were then cast into a disk--shaped mould with diameter of 8mm and height of 1mm. PLLA disks were formed by evaporating the solvent overnight. POMaC disks were formed by UV--exposure at an intensity of 10mJ/cm 2 s for 4min. Both PLLA disks and POMaC disks were sterilized in 70% ethanol and repeatedly washed in ethanol and deionized water prior to implantation. PLLA and POMaC disks were subcutaneously implanted for 5 weeks in adult male Lewis rats. Explanted tissues were formalin fixed, paraffin embedded, sectioned and stained for H&E, Masson's Trichrome, CD3, CD68, and CD163 at the Pathology Research Program (PRP) at the University Health Network. Tissue sections were scanned and cells within 200μm (400pixels) from the implant surface were quantified for positive staining of Masson's Trichrome, CD3, CD68, or CD163.
Quantification of cell density in AngioChip tissues. Cell density was quantified from histology sections based on previously described stereology method 3 according to the following formula: (Eq. 1) where N V is the number of nuclei per unit volume, N A is number of nuclei per unit area on a histology section, t is the section thickness corrected for tissue shrinkage, and D is the average diameter of cell nuclei. First, the average cell nuclei diameter was determined from confocal images. The average diameter of at least 40 nuclei was used. The number of nuclei per unit area on the randomly selected areas of the cross--section of the AngioChip tissues was determined from 3μm thick histology sections. The micro--channel area was excluded from the area calculation. A correction factor of 5.7% reported previously 3 was used to account for tissue shrinkage during tissue preparation for histology. For hepatic tissues, density was quantified for the AngioChips containing hESC--hepatocytes and 20% of human MSCs in the parenchyma. For cardiac tissues, density was quantified for the AngioChips based on hESC--derived CMs.
Histology and immunofluorescent staining. Immuno--fluorescent staining was performed to assess the morphology of cultivated AngioChip tissues. The AngioChip cardiac tissues and the Biowire or monolayer controls were first fixed in 4% (w/v) paraformaldehyde in PBS for 15min at room temperature. Then, the tissues were permeated in situ and blocked with 5% FBS and 0.25% Triton×100 in PBS for 1 hour. Next, the AngioChip cardiac tissues were incubated in primary antibody, sarcomeric α--actinin (Mouse, 1:200 dilution, Sigma) overnight at 4˚C, followed by incubation with corresponding secondary antibodies, Alexa 488 anti--mouse IgG (1:200 dilution, Sigma) or TRITC anti--mouse (1:200 dilution, Sigma) and F--actin (Phalloidin 660 conjugated, Sigma) for 1hour. The cardiac tissues were imaged with Olympus FV5--PSU confocal microscope with 10× and 40× objectives. The AngioChip hepatic tissues and the collagen sandwich controls were first fixed in 4% (w/v) paraformaldehyde in PBS for 15min at room temperature. Then the tissues were permeated in cold methanol on ice for 2min. Next, the tissues were incubated in primary antibody, E--cadherin (Mouse, 1:200 dilution, BD Biosciences) and Albumin (Goat, 1:200 dilution, Bethyl) overnight at 4˚C, followed by incubation with corresponding secondary antibodies, Alexa 647 anti--mouse IgG (1:200 dilution, Sigma) and TRITC anti--goat (1:200 dilution). Prior to fixation, some hepatic tissues were incubated with Carboxy--DCFDA (5µM, Life Technologies) in HBSS (with Ca 2+ /Mg 2+ ) for 20min as described to visualize bile canaliculi 6 . Carboxy--DCFDA is a substrate of MRP2, a hepatocyte efflux transporter, thus it can be used to observe the canalicular uptake and efflux, as well as to label the bile canaliculi structures, crucial in the transport of bile acid from hepatocytes to bile ducts 6 . To visualize the vasculature, endothelialized AngioChip scaffolds were stained with CD31 (Mouse, 1:200 dilution, Sigma), followed by Alexa 647 anti--mouse IgG (1:200 dilution, Sigma) as described previously 4 . Live and dead staining was performed with carboxyfluorescein diacetate (CFDA, 1:1000 dilution, Invitrogen) and propidium iodide (PI, Invitrogen) in PBS at 37˚C for 30min. To visualize the cross--section of the cardiac tissue with live and dead staining, the tissues were sliced in half transversely and rotated manually to show the cross--section. To visualize both endothelial cells and parenchymal cells, cultivated tissues were fixed in 10% formalin for 3 days at 4˚C, paraffin embedded, and sectioned into 3μm slices for histology at the Pathology Research Program Laboratory of Toronto University Health Network. Histology sections show the tissue cross--section in the transverse direction and were stained with Hematoxylin and Eosin (H&E), Masson's Trichrome stain, CD31, or albumin.
Urea assay on hepatic tissue: The AngioChip hepatic tissues were incubated with 1mL hepatocyte media, containing 10mM ammonium bicarbonate, in the main wells and perfused with 1mL endothelial cell media, containing 10mM ammonium bicarbonate, from the inlet wells. After 24hr incubation, the media from the main wells and the outlet wells were collected. The media were briefly centrifuged (300×g for 5min) to remove any cells and frozen at --20˚C. Urea was detected using QuantiChrom Urea assay kit (BioAssay Systems) as per the manufacturer's instructions. The media were typically used undiluted for detection. The media baseline reading for urea was determined from media collected at the same perfusion condition but without the presence of liver cells in the AngioChip scaffold. To quantify cell number of AngioChip tissue and collagen sandwich culture, the tissues were first ground and the cells were lysed in cell lysis buffer (Cell Signaling). The supernatant was collected and analyzed with Quant--iT™ PicoGreen® dsDNA Assay Kit according to manufacturer's instruction and correlated to a standard curve.
Liver drug test. On day 3 of culture, terfenadine (10µM, Sigma) in hepatocyte culture media was placed in the inlet wells and perfused through the endothelialized hepatic tissue at 0.7μL/min (0.62dynes/cm 2 , Re, 0.01) for 24hr. After 24hr incubation, the media in the inlet wells, middle wells, and outlet wells were collected and analyzed for the concentration of fexofenadine with Liquid Chromatography-Mass Spectrometry (LC--MS) and correlated to a standard curve. DMSO was used to dissolve the drugs initially. The final concentration of DMSO in the media was always less than 0.1%.
Lactate dehydrogenase (LDH) assay on cardiac tissue: The AngioChip cardiac tissues were cultured with cardiomyocyte media with or without media perfusion through the built--in network at 0.7μL/min (0.62dynes/cm 2 , Re, 0.01). From the middle wells where the tissues resided, 1mL of cardiomyocyte media was collected and replaced with new media every day for 6 days. The collected media were analyzed for lactate dehydrogenase (LDH) concentration with a LDH toxicity assay kit (Cayman Chemical) as per manufacturer's instructions and correlated to a standard curve.
Cardiac drug test. On day 7 of culture, the spontaneous contractions of the cardiac tissues were recorded as brightfield videos. Then, epinephrine (10µM, Sigma) or digoxin (10µM, Sigma) in cardiac culture media were placed in the inlet wells and perfused at 0.7μL/min (0.62dynes/cm 2 , Re, 0.01) through the endothelialized cardiac tissue for 30min. After 30min incubation, the spontaneous contractions of the cardiac tissues were recorded again as brightfield videos. The frequency of contraction was analyzed from the recorded videos with Image J. DMSO was used to dissolve the drugs. The final concentration of DMSO in the culture media was diluted to less than 0.1% (v/v).
Functional characterization of engineered cardiac tissue. To stimulate the AngioChip cardiac tissue and measure their electrical excitability parameters, two carbon electrodes spaced 1cm apart were placed within each main well on the opposite sides of the tissue in parallel. The electrodes were connected to an external electric stimulator (Grass s88x) with platinum wires. Using monophasic pulses of 2ms duration and frequency of 1 pulse per second, the excitation threshold (minimum voltage at which synchronous contractions of 75% of the tissue in the field of view can be observed) was first determined. Then the maximum capture rate (maximum beating frequency) was determined at 200% of the determined excitation threshold voltage. The amplitude of the tissue contraction was determined from the change of width in the engineered tissue between contractions normalized to the tissue width at the relaxed state. To track the process of tissue remodeling, a bright--field image of the AngioChip cardiac tissue was taken daily in the first 5 days after cell seeding (Olympus SC30 camera). For the purpose of comparison with other published studies and to quantify the gel compaction rate for the rat and human cardiac constructs, the AngioChip tissue area was measured from brightfield images at different time points. Conduction velocity was assessed upon staining the tissues with a voltage sensitive dye, Di--4--ANNEPS, as we described previously 5 .

POMaC degradation properties, mechanical properties, and rationale for selection
Degradation in PBS is commonly employed to measure the degradation rates of hydrolytically degradable polymers, such as POMaC 7 , in vitro. Accelerated degradation at high pH is usually performed to confirm that the bonds that participate in hydrolysis are accessible within the biomaterial. The AngioChip scaffolds degraded gradually in phosphate buffered saline (PBS) over months and to completion in 3 days under alkaline conditions (0.1M NaOH) (Supplemental Figure 8a,b), time--frames that were consistent with those reported previously for the base material used in AngioChip production, POMaC 7 .
By varying the UV energy (factor A), heat exposure time (factor B), monomer ratio (factor C) and porogen concentration (factor D), we were able to vary the POMaC bulk polymer elasticity over a wide range (Supplementary Table 3) and come up with an empirical equation, using a design of experiment (DOE) approach, that enables us to fine--tune the mechanical properties for future applications.
where only the significant effects (i.e. p < 0.05) were included. In Eq. 3, b n coefficient subscripts are labelled as follows: o=intercept of the model, A=UV energy, B=heat duration, BD=interaction between heat duration and porogen content, CD=interaction between monomer ratio and porogen content, ABD=interaction between UV energy, heat direction, and porogen content and ε is the residual. The coefficients for the partial model are shown in Supplementary Table 4.

Supplementary Table 4: Coefficients for the DOE model
The results demonstrated that the Young's modulus directly depended on the UV exposure and the thermal crosslinking times, while porogen content exhibited interactive effects with both heat exposure time and monomer ratio. The model also detected an interactive effect of three independent factors: UV crosslinking time, heat exposure and porogen content.
Trends for the main factor effects were as expected. Over the ranges investigated, UV energy had the POMaC was therefore selected due to elasticities, ranging from 53--1423kPa, favorable for soft tissue applications (Supplementary Figure 8c, Supplementary Table 3) and documented blood compatibility of citric acid--based elastomers 8,9 . FDA approved polyesters poly(lactic acid), poly(glycolic acid), poly(ε--caprolactone) and their co--polymers are significantly more stiff (Young's moduli range from 0.1GPa to 10 GPa 10 ) and brittle, thus they are not suitable for AngioChip production for the purpose of soft tissue engineering. Additionally, they are not suitable for manufacturing of blood contacting devices and vascular tissue engineering as they are thrombogenic 11 without additional modification. Here, we demonstrated AngioChip production from POMaC, but we propose that other scaffold materials might be used in the future as long as they satisfy the general criteria of tissue specific elasticity and non--thrombogenicity.

Culture medium distribution in AngioChip scaffold compartments
To characterize culture medium flow in different compartments, 4mL of PBS was loaded into the inlet well and 1mL in the main well of the multi--well plate, where an AngioChip scaffold with 10µm holes was placed. The PBS was allowed to perfuse through the AngioChip for a total of 24hr, at which point the PBS was collected from the inlet, main and the outlet wells for the purpose of closing the mass balance on the perfused fluid. As illustrated in the Supplementary Figure 8d, most of the perfused fluid ended up in the outlet well as expected. Specifically at 24hr, there was 2.7±0.2mL in the inlet well, 1.2±0.1mL in the main well and 1.1±0.1mL in the outlet well, indicating there was no significant leakage of the fluid into the tissue space, even without the presence of endothelial cells in the AngioChip lumen.

Diffusivity in AngioChip parenchymal space
We observed transient FITC diffusion from the outer AngioChip channel lumen through the parenchymal layer and into the media reservoir, which behaved as an infinite sink since the volume of the media was 67 times greater than the AngioChip volume. Fluorescence intensity profiles were obtained using Image J at the mid--line of the AngioChip, perpendicular to the channel wall, 20min and 30min after initiation of FITC perfusion (10µM in PBS). Focusing on the diffusion from the outer channels, it was possible to apply the theory for transient diffusion in the semi--infinite media, as described 12 . Using this approach the diffusion coefficient for FITC in the parenchymal space was estimated to be 3.4--3.8×10 --11 m 2 /s, only slightly lower than the diffusion coefficient of FITC in PBS as reported previously (2.7×10 −10 m 2 /s) 13 , indicating that the polymer lattice in the parenchymal layer did not dramatically hinder solute diffusion.

Rationale for selection of micro--hole sizes
The micro--hole sizes of 10µm and 20µm were selected to enable cell extravasation from the main AngioChip vasculature into the parenchymal space. The micro--holes of 10µm in diameter and 50µm in spacing were selected to enable monocyte extravasation based on their size and a previously published microfluidic lung--on--a--chip 14 that included a membrane with 10µm micro--holes spaced 40--50µm apart, which allowed robust transmigration of monocytes (12µm in diameter for THP--1 monocytes 15 ) in previous studies 14 as well as here (Figure 3 m--r). Micro--holes of 20µm in diameter were selected to allow robust endothelial sprouting since small capillaries can range from 5--10µm in diameter. We found that within micro--holes of 20µm in diameter, multiple endothelial cells could wrap into a tubular structure (Figure 3e--g).

Polymer biocompatibility
At 5 weeks post sub--cutaneous implantation in a rat model, both poly--L--lactic acid (PLLA) and POMaC scaffolds were visible at the implantation site, indicating POMaC did not biodegrade fully over 5 weeks. Qualitatively, the appearance of histology sections of PLLA and POMaC from the biocompatibility study was similar (Supplementary Figure 26). POMaC exhibited less fibrosis in comparison to PLLA, as assessed by collagen coverage from Mason's Trichrome staining (Supplementary Figure 26). Quantification of immunostaining for macrophage and T--cell specific markers demonstrated that although there were more M2, pro--healing, macrophages with the POMaC scaffold implantation, there were also more T--cells in this group compared to the PLLA group. The ratio of pro--healing to total macrophages was similar in the PLLA and POMaC groups. Further studies are necessary to delineate similarities between PLLA and POMaC responses at different times post implantation, to evaluate the contribution of different scaffold elasticities to the observed responses, as well as to determine the time required for the complete degradation of both scaffolds in this animal model.

Cell culture considerations for construction of mm--thick tissues
We made two simple modifications to our culture procedure developed for the single--layer AngioChip tissues in order to improve both cell density and organization in thick multi--layer AngioChip tissues: 1) We used centrifugation assisted cell seeding to seed cells uniformly in the entire volume of a 3--layer AngioChip scaffold and 2) We extended the liquid reservoir in the AngioChip bioreactor to achieve a higher pressure head and a higher perfusion rate (11.7 ± 3.6 mL/day, n=4). For seeding of thin tissues, we used a simple pipetting for cell seeding and lower pressure heads (i.e. perfusion rates, 1.1 ± 0.1mL/day) that were not sufficient to support cell growth in the thick multi--layer scaffolds. We also waited for a minimum of 3 days in perfusion culture to assess cell distribution and organization in these high--density constructs (Figure 5q--w, Supplementary Figure  19, 20). Our previous work conclusively showed that in the thick high--density tissues with insufficient oxygen and nutrient supply (i.e. tissues cultured under static conditions), massive cell death occurred within a period of hours (e.g. 4.5 hr 16 ) after seeding of metabolically active cells such as cardiomyocytes, resulting in the inability of the cells in the scaffold interior to elongate and a non--uniform cell distribution (i.e. a void interior) 16,17 . Centrifugation assisted seeding is critical for uniform initial cell distribution in thick scaffolds and providing sufficient pressure head is critical for proper tissue perfusion. In principle, these factors can be tuned for tissues of thickness greater than 2mm as described here using scalable AngioChip--based techniques that rely on multi--layer scaffolds with one inlet and one outlet, or on the stacking of individual single--layer AngioChip tissues to make a thicker structure. The scaffold was perfused with barium sulfate solution through its internal network hence increasing its density for improved visualization. The thickness of the scaffold network wall was 50μm. The inlet, outlet, and the first order branch had an inner luminal dimension of 50μm by 200μm. The second order branch had an inner luminal dimension of 50μm by 100μm. The network was designed so that the endothelial cells in the first and second order branches experienced the same level of shear stress. The networks on each layer were connected through a vertical channel and were 300μm apart in z--axis. The scaffold mesh was made of 50μm struts. The struts were spaced 250μm apart in the long--edge direction, 100μm apart in the short--edge direction, and 50μm apart in the z--axis. This version of the multi--layer scaffold has thickness of 500μm.  One configuration is artery bypass where the inlet and outlet of the AngioChip network were both connected to an artery. The blood from the femoral artery was routed through the AngioChip and back to the artery. Another is artery to vein loop where the inlet of the AngioChip network was connected to an artery while the outlet was connected to a vein. The blood from the femoral artery was routed through the AngioChip and delivered directly to the femoral vein, going back to the heart. Surgical vessel clamps were used to shut the artery and vein temporarily during surgery. After the surgical procedure, vessel clamps were removed and blood was perfused immediately through the AngioChip network in both surgical configurations. Vessel pulsation can also be observed on the AngioChip network.