CROSS-REFERENCE TO RELATED APPLICATION

[0001] This application claims priority to U.S. Provisional Application entitled “CULTURE CHAMBER SYSTEM FOR CARDIAC TISSUE,” having serial number 62/796,868, filed on January 25, 2019, which is entirely incorporated herein by reference.

GOVERMENT LICENSE RIGHTS

[0002] This invention was made with government support under grant number 1 1675980 awarded by the National Institutes of Health (NIH). The government has certain rights in the invention.

SEQUENCE LISTING

[0003] This application contains a sequence listing filed in electronic form as an ASCII.txt file entitled“(222104-2920) Sequence Listing_ST25.txt”, created on January 24, 2020. The content of the sequence listing is incorporated herein in its entirety.

BACKGROUND

[0004] Cell culture is a process through which cells are grown under controlled conditions outside their natural environment. For example, a sample of cells can be isolated from living tissue and maintained in an environment suitable for culturing or growing the cells. The conditions in the environment may vary depending upon the type of cell or cells being cultured. The environment may be contained and maintained in a type of vessel or chamber with a substrate or medium that supplies nutrients (e.g. , amino acids, carbohydrates, vitamins, minerals), gases (e.g. , CO ₂ , O ₂ ), and regulates the chemical environment (e.g., pH, pressure, temperature). Some cells may need an artificial substrate to grow upon while others can be grown suspended in a culture medium. The term“cell culture,” today, often refers to culturing cells derived from multi-cellular eukaryotes, such as animal and human cells, in contrast with fungal, microbiological, and other single-cellular cultures.

SUMMARY

[0005] In certain aspects, described herein are culture chamber systems for cardiac or other mammalian cells sensitive to electrical and/or mechanical stimulation. In an embodiment according to the present disclosure, a culture chamber system as described herein comprises a culture chamber, the culture chamber comprising a well, a flexible membrane enclosing a side of the well, and a pair of posts extending within the well and each adhered at one end to the flexible membrane; a flow loop to supply media to the well in a first phase of a cycle, wherein the flexible membrane distends due to a weight of the media in the well during the first phase of the cycle; a support base to buttress a distention of the flexible membrane due to the weight of the media in the well; and a pump to oppose the distention of the flexible membrane during a second phase of the cycle. The well can be formed in a block of silicone.

[0006] The flow loop can comprise a media reservoir to hold a volume of the media; an outlet tube for the media to flow from the media reservoir to the well; and an inlet tube for the media to flow from the well back to the media reservoir. In further embodiments according to the present disclosure, a flow loop can further comprise: a first one-way valve in-line with the outlet tube; a second one-way valve in-line with the inlet tube; and a variable flow constrictor in-line with the outlet tube, wherein the first one-way valve and the second one way valve permit flow of the media in one direction through the outlet tube and the inlet tube.

[0007] In embodiments according to the present disclosure, culture systems as described herein can further comprise an electrical actuator configured to provide an electrical stimulus within the well during the second phase of the cycle.

[0008] Also described herein are methods of operating a culture chamber system as described herein. In an embodiment according to the present disclosure, a method of operating a culture chamber system comprises supplying media to a well of a culture chamber in a first phase of a cycle, wherein a flexible membrane of the culture chamber distends due to a weight of the media in the well during the first phase of the cycle; and pumping a fluid to the culture chamber to oppose distention of the flexible membrane during a second phase of the cycle. Methods as described herein can further comprise providing an electrical stimulus within the well during the second phase of the cycle.

[0009] The culture chamber of methods as described herein can comprise the well; the flexible membrane enclosing a side of the well; and a pair of posts extending within the well and each adhered at one end to the flexible membrane. In embodiments according to the present disclosure, the culture chamber system of methods as described herein can further comprise a flow loop for suppling the media to the well in the first phase of a cycle; a support base to buttress a distention of the flexible membrane due to the weight of the media in the well during the first phase of the cycle; and a pump for pumping the fluid to oppose the distention of the flexible membrane during the second phase of the cycle.

[0010] Described herein are bioreactors. In an embodiment according to the present disclosure, a bioreactor can comprise a culture chamber, the culture chamber comprising a well; a flexible membrane enclosing a side of the well; and a pair of posts extending within the well and each adhered at one end to the flexible membrane. In embodiments according to the present disclosure, the well can be formed in a block of silicone or polydimethylsiloxane (PDMS). In embodiments according to the present disclosure, the well can further comprise a fluid inlet and outlet. In embodiments according to the present disclosure, the flexible membrane can be about 80 pm thick. The thickness is a dimension from a side of the flexible membrane facing an opening of the wall to the opposite side of the well abutting and/or facing a support base. In embodiments according to the present disclosure, the flexible membrane can be polydimethylsiloxane (PDMS). In embodiments according to the present disclosure, the pair of posts can be polydimethylsiloxane (PDMS). In embodiments of the present disclosure, bioreactors can further comprise a support base beneath the well configured to buttress distention of the flexible membrane. In embodiments of the present disclosure, bioreactors can further comprise an electrode on a side of one or both of the pair of posts. In embodiments of the present disclosure, bioreactors can further comprise one or more electrodes on one or more sides of the well opposite the pair of posts. In embodiments of the present disclosure, the well of bioreactors as described herein is elliptical in shape. In embodiments of the present disclosure, the well of bioreactors as described herein can further comprise 2% agarose or other culture material (culture materials that can coat the well are known in the art and can be varied by the skilled artisan according to the specific type of cell to be cultured in the well

[0011] Also described herein are methods of culture induced-pluripotent stem cell- derived cardio myocytes (iPS-CM) or progenitors thereof. In an embodiment, a method of culturing induced-pluripotent stem cell-derived cardio myocytes, comprises providing a culture chamber system as described herein; plating fibrin-encapsulated iPS-CM progenitor cells in the well of the culture chamber; culturing the fibrin-encapsulated iPS-CM progenitor cells for five days following plating; supplying media to a well of a culture chamber in a first phase of a cycle after five days, wherein a flexible membrane of the culture chamber distends due to a weight of the media in the well during the first phase of the cycle; and pumping a fluid to the culture chamber to oppose distention of the flexible membrane during a second phase of the cycle. In embodiments, methods as described herein can further comprise providing an electrical stimulus within the well during the second phase of the cycle. In embodiments of methods as described herein, the culture chamber comprises a bioreactor as described herein. In embodiments of the present disclosure, the culture chamber system of methods as described herein comprises a flow loop for suppling the media to the well in the first phase of a cycle; a support base to buttress a distention of the flexible membrane due to the weight of the media in the well during the first phase of the cycle; and a pump for pumping the fluid to oppose the distention of the flexible membrane during the second phase of the cycle. In embodiments of methods according to the present disclosure, methods further comprise providing to the cells in the culture chamber after five days a peak systolic pressure of 5 mm Hg, an end-diastolic pressure of 1 mm Hg, and a maximum strain of 0.2% for a first period of 24 hours (1 day). In embodiments of methods according to the present disclosure, methods further comprise increasing the pressure and maximum strain after the first period of 24 hours to 10 mm Hg, 3 mm Hg, and 0.7%, respectively, for a second period of 24 hours (1 day). In embodiments of methods according to the present disclosure, methods further comprise increasing the pressure and maximum strain after the second period of 24 hours to 15 mm Hg, 5 mm Hg, and 1.2%, respectively, for a third period of 24 hours (1 day). In embodiments of methods according to the present disclosure, methods further comprise increasing the pressure and maximum strain after the third period of 24 hours to 30 mm Hg, 10 mm Hg, and 2.3%, respectively, for a fourth period of 96 hours (4 days). In embodiments of methods according to the present disclosure, methods further comprise maintaining the cycle frequency at 1 Hz.

BRIEF DESCRIPTION OF THE DRAWINGS

[0012] Aspects of the present disclosure can be better understood with reference to the following drawings. It is noted that the elements in the drawings are not necessarily to scale, with emphasis instead being placed upon clearly illustrating the principles of the embodiments.

[0013] FIG. 1 illustrates an example culture chamber system according to various embodiments described herein.

[0014] FIGS. 2A and 2B are representative photographs of the culture chamber of the system illustrated in FIG. 1 according to various embodiments described herein. [0015] FIG. 3 illustrates a diastolic phase of a pumping process performed using the culture chamber system illustrated in FIG. 1 according to various embodiments described herein.

[0016] FIG. 4 illustrates a systolic phase of a pumping process performed using the culture chamber system illustrated in FIG. 1 according to various embodiments described herein.

[0017] FIGs. 5A-5E are photographs showing reduced to practice embodiments of the present disclosure and a diagram of a method according to the present disclosure. FIG. 5A shows biomimetic cardiac tissue model (BCTM) cell culture chamber with agarose mold prior to the addition of cells within the fibrin gel. FIG. 5B shows BCTM cell culture chamber with the cell-laden fibrin gel suspended between the 2 posts. This image was taken 2 days after removing the agarose mold. The gel originally takes the shape of the agarose mold but contracts over time as shown in the image. FIG. 5C shows the position of the posts within the cell culture chamber at the end of systole. The fiber experiences no strain at this point as the posts are at the initial resting position. FIG. 5D shows the position of posts within the cell culture chamber at the end of diastole. The thin membrane contacts with the insert below and deforms along its surface, thus pulling the posts away from each other and applying uniaxial strain to the fiber. FIG. 5E is a diagram representing how the BCTM reproduces the cardiac cycle. The BCTM imposes uniaxial strain onto the cells by using a flexible membrane that is stretched over an arched surface placed below the membrane. Yellow indicates the cell laden fibrin fiber.

[0018] FIG. 6 is a graph showing pressure as a function of maximum strain. Pressure-volume loops that human induced-pluripotent stem cells cardiomyoctytes (hiPSC- CMs or hiPS-CMs) were subjected to during the course of the 7-day study. The systolic/diastolic ratio was maintained at 40/60% and the frequency was maintained at 1 Hz throughout the duration of the experiment.

[0019] FIGs. 7A-7B are fluorescent micrographs showing Static Controls (FIG. 7A) and BCTM Stimulated samples (FIG. 7B) with the entire fiber stained with cardiac troponin T (cTnT), alpha sarcomeric actin (a-SA), and 4',6-diamidino-2-phenylindole (DAPI) with magnified images of a-SA (bottom), cTnT (middle), and combined image (top row).

[0020] FIGs. 8A-8B are groups of micrographs of static controls (FIG. 8A; left side) and stimulated samples (FIG. 8B; right side). (Top) Visualization of extracellular matrix (ECM). (Middle) Representative images of sarcomeres and associated structures within cardiomyocytes. Arrows indicate lipid deposits (black) and an example of unorganized sarcomeres (white) both of which were common throughout the static controls. (Bottom) High magnification images of representative sarcomeres from the static controls and stimulated samples. The black arrow on the stimulated sample marks a white area along the middle of the sarcomere that was found consistently in the stimulated sarcomeres and what could be the formation of the M-line.

[0021] FIGs. 9A-9G are graphs of quantitative real-time polymerase chain reaction (qRT-PCR) results of static controls vs. BCTM stimulated samples normalized to GAPDH expression levels. Statistically significant upregulation was observed in genes GAT4, MYL2 , CAM2KB , and MYH7 in stimulated samples (N=5) vs. static samples (N=4) as denoted with asterisks.

[0022] FIG. 10 is a graph of data collected from videos made of the fibers undergoing electrical stimulation, showing fractional shortening (%) as a function of a pacing frequency (Hz), at frequencies: 1 Hz, 1.3 Hz, 1 .5 Hz, and 2.0 Hz. Videos were also made of the intrinsic (spontaneous) fractional shortening of each fiber (N=4). The stimulated samples showed significantly higher fractional shortening at 1 .3, 1 .5, and 2.0 Hz. The intrinsic contractions were not significantly different.

[0023] FIG. 1 1 is an illustration of cardiac pressure and volume overload. Pressure overload is associated with an increase in end systolic pressure that causes an increase in the ventricular wall thickness through myocyte thickening caused by sarcomere addition. Volume overload occurs when the end diastolic volume is increased leading to a constant increase in residual ventricular volume that over time causes wall thinning through myocyte thinning and elongation.

[0024] FIGs. 12A-12C are re duced-to- practice images of an embodiment of a cell culture chamber according to the present disclosure. FIG. 12A is an image taken 2 days after removing the agarose mold and removing the cell laden fibrin gel from the two posts. The gel originally forms around the posts and takes the shape of the agarose mold but contracts over time. Once the gel contracts away from the agarose mold it becomes a free-standing fiber suspended between two posts. FIG. 12B shows the position of the posts within the cell culture chamber at the end of systole. The fiber experiences no strain at this point as the posts are at the initial resting position. FIG. 12C shows the position of posts within the cell culture chamber at the end of diastole. The thin membrane contacts with the insert below and deforms along its surface thus pulling the posts away from each other and applying uniaxial strain to the fiber (scale = 4 mm).

[0025] FIGs. 13A-13C are graphs of H9c2 pressure-volume loops. Plots of the pressure vs. percentage strain that H9c2 cells were subjected to within the Cardiac Tissue Chips (CTC) during the course of an experiment. The pressure and percentage strain were gradually increased over the course of 1 hour until the final pressure and strain values shown in the plots were obtained and maintained at that level until the conclusion of the experiment. Percent strain is shown in normal conditions (FIG. 13A), volume overload conditions (FIG. 13B), and pressure overload conditions (FIG. 13C).

[0026] FIGs. 14A-14D are micrographs of stained tissue showing gross morphology and alignment. Hematoxylin and eosin (H&E) staining of paraffin embedded tissue sections show alignment and uniform distribution of H9c2 cells within the fibre regardless of location of the fibre. The image[s] shown (FIG. 14A and insets FIGs. 14B-14D) is from a sample subjected to normal pressure-volume loading but the alignment and organization are representative of all samples including static, pressure overload, and volume overload samples (scale: top = 500 pm, bottom = 50 pm).

[0027] FIGs. 15A-15H are micrographs of Masson’s trichrome staining. Paraffin embedded and sectioned H9c2 fibres were imaged following Masson’s trichrome staining. FIGs. 15A-15B (Static) show unstimulated tissue fibre that was maintained in static culture. FIGs. 15C-15H (Normal, Pressure, and, Volume) show tissue fibres stimulated within the CTC under normal, pressure overload, and volume overload conditions respectively. The pressure overload sample (FIGs. 15E-F) shows significant increases in fibrosis along the centre line of the fibre and is clearly thicker than the other samples. Thinning is also noted in the volume overload sample (FIG. 15G-15H). There were minimal levels of fibrosis within the static, normal, and volume overload fibres (scale: left = 100 pm, right = 50 pm).

[0028] FIG. 16 shows aspect ratio measurements. Static, normal, volume overload, and pressure overload samples labelled ‘static’, ‘normal’, ‘pressure’, and ‘volume’ respectively (N=5 for each group). The volume overload showed a significant increase in length and is clearly elongated in the image so much so that it has buckled once it returned to its original resting state. The increase in length resulted in a significant decrease in aspect ratio compared to static, normal, and pressure overload samples. Pressure overload samples significantly increased in width but not length, resulting in a significant increase in aspect ratio when compared to static, normal, and volume overload samples. The difference between normal and static samples was not significant although a decrease in aspect ratio was found in the normal samples (scale = 4 mm).

[0029] FIGs. 17A-17N are gene expression profiling graphs showing relative expression of genes associated with: fibrosis (FIGs. 17A-17D); matrix remodeling (FIGs. 17E-17H); cytoskeleton (FIGs. 17I-17K); and oxidative stress (FIGs. 17L-17N). The results are from qRT-PCR array on H9c2 cells stimulated for 48 hours under static (N=3), normal (N=3), volume overload (N=3), and pressure overload (N=3) hemodynamics. Genes are classified into 4 categories, ‘Fibrosis’ (FIGs. 17A-17D), ‘Extracellular matrix remodelling’ (FIGs. 17E-17H), ‘Cytoskeleton’ (FIGs. 17I-17K), and‘Antioxidants’ (FIGs. 17L-17N). All genes were normalized to GAPDH levels.

[0030] FIG. 18 is an illustration of a system according to the present disclosure.

[0031] FIG. 19 is an embodiment of a mold according to the present disclosure (FIG.

12A), shown in position with the posts of the cell culture chamber well.

[0032] FIG. 20 shows an embodiment of a culture well as described herein.

[0033] FIGs. 21A-21 C show an embodiment of the BCTM with integrated electrodes (FIG. 21 A). FIG. 21 B shows a side view, and FIG. 21 C are graphs showing synchronized delivery of electrical stimulation with the onset of systole within the BCTM.

[0034] FIG. 22 is an illustration showing hiPS-CM differentiation vs. Heart Development: Contrasting hiPS-CM differentiation with embryonic heart development with focus on hemodynamic loading seen during different phases of heart development. Embryonic heart development is split into 3 phases (green, red and blue).

[0035] FIGs. 23A-23E: Masson Trichrome staining (FIG. 23A) Normal, (FIG. 23B) Pressure overload conditions. (B) Cardiac tissue constructs maintained under normal pressure-volume loading (FIG. 23C) and static culture with TGF- b (FIG. 23D) than under conditions of pressure overload with TGF- b (FIG. 23E).

[0036] FIG. 24A and 24B show an embodiment of a reduced-to-practice embodiment of the present disclosure used to recreate left ventricular assist device (LVAD) support within the BCTM (FIG. 23A) and pressure-volume loops showing changes in preload (left ventricular volume) and afterload (pressure) under normal, heart failure and LVAD support conditions (FIG. 24B).

[0037] FIGs. 25A-25C: FIG. 25A is a reduced-to-practice embodiment of a complete BCTM flow loop to subject engineered cardiac tissue to pressure-volume loading; FIG. 25B is a schematic and actual image of the 3D muscle strip integrated within the BCTM flow loop chamber where pressure-volume changes can be imposed. FIG. 25C are graphs showing examples of putative changes from normal pressure-volume loading to pathological conditions like pressure and volume overload.

DETAILED DESCRIPTION

[0038] As noted above, cell culturing is a process through which cells are grown under controlled conditions. In that context, a new culture chamber system is described herein. In one example, the culture chamber system includes a culture chamber, where the culture chamber includes a well, a flexible membrane enclosing the bottom side of the well, and a pair of posts extending within the well and each adhered at one end to the flexible membrane. Cells, such as cardiac cells, are grown into a tissue between the posts. The culture chamber system is designed in part to mimic the 4 phases of the ventricular cycle. Initially, the ventricle is in a relaxed state following which the filling of the ventricle begins. The relaxed ventricle gradually fills resulting in stretching of the ventricular walls (e.g., the cardiac tissue). The magnitude of stretch increases until a maximum value is attained at the end of the filling phase. Once the filling phase is complete, the contractions begin. During the contraction phase, the cardiac tissue contracts resulting in isovolumetric compression of fluid within the culture chamber. The pressure within the ventricular chamber increases until the pressure is sufficient to overcome the fluidic resistance (e.g. , until the pressure within the ventricular chamber is greater than the pressure outside the chamber). Once the pressure within the chamber is greater than the outside pressure, the ejection phase begins. The fluid is ejected into the outlet (aorta) and the filled volume is ejected. Then, the relaxation phase begins. The contracted ventricular walls relax and the pressure within the ventricular chamber return to the initial value (zero, in the healthy heart) and then the filling begins again followed by contractions and relaxation. The relaxation and filling phases together represent the Diastole or the Diastolic Phase referred to in subsequent discussions as the“first phase,” and the contraction and ejection phases represent Systole or the Systolic Phase referred to in subsequent discussions as“second phase”. The culture chamber system further includes a flow loop to supply media to the well in a first phase of a cycle, and the flexible membrane distends due to a weight of the media in the well in the first phase of the cycle. The culture chamber system further includes a support base to buttress a distention of the flexible silicone membrane due to the weight of the media in the well, and a pump to oppose the distention of the flexible silicone membrane during a second phase of the cycle. The culture chamber system can reproduce the effects of electrical and pressure-volume changes (e.g., pressure and stretch) seen in any chamber of the heart at any stage of development. The electrical stimulation can be precisely coordinated with mechanical loading (e.g., at the onset of the systole phase to represent normal cardiac function or mismatched to represent disease) to ensure that the electrical impulse is coordinated with cardiomyocyte contractions. Using the culture chamber system, both electrical and mechanical stimulation can be applied to live, three-dimensional cardiac tissue constructs that mimic the cellular compositions in embryonic, fetal, neonatal, and adult cardiac tissue.

[0042] Further, various pathological conditions including pressure overload, volume overload, and heart failure can be accurately replicated using the culture chamber system and the cardiac tissue grown within the culture chamber system. The cardiac tissue grown within the culture chamber system can also be seamlessly integrated with tissues of other organs to create a more complex body-on-a-chip platform, where the cardiac tissue chamber acts as a pump to ensure perfusion within the system.

[0043] Additionally, various time-of-day dependent changes in blood pressure, stretch, and heart rate can be reproduced using the culture chamber system. The cardiac tissue grown in the culture chamber system can be used to recreate time-of-day dependent pressure-volume changes (e.g., activity and rest changes) and pressure-volume changes associated with exercise (e.g. , increased heart rate). Mechanical unloading, as seen with mechanical circulatory support (MCS) and/or left ventricular assist devices (LVADs), can also be reproduced.

[0044] The cardiac tissues developed using the culture chamber system can be used for various research and development purposes. For example, the cardiac tissues can be used as a platform to generate mature cardio myocytes (human and other species) by carefully reproducing the electrical and mechanical stimulus associated with embryonic heart development, to produce structurally and functionally mature cardio myocytes for research and clinical applications (e.g. , for drug testing, drug discovery, and regenerative medicine). The cardiac tissues can also be used to create models of cardiovascular disease, including conditions like heart failure, pressure overload, and volume overload. The cardiac tissues can also be used to create disease models where interventions (pharmaceutical and device) can be tested. Additionally, the cardiac tissues can be used as a physiologically relevant model to test drug induced cardiotoxicity of any pharmaceutical drug (e.g., off target effects on cardiomyocytes particularly arrhythmogenic potential). [0045] FIG. 1 illustrates an example culture chamber system according to various embodiments described herein. The culture chamber system is provided as a representative example of a system including a culture chamber for cardiac tissue. According to the examples described herein, the system can be relied upon to culture or grow cells that form cardiac (e.g., heart) tissue, although the system can be used to culture cells other than those that form cardiac tissue. The system is not drawn to scale in FIG. 1 , can include other components not illustrated in FIG. 1 , and may omit one or more of the components illustrated in FIG. 1 , in various embodiments.

[0046] Among other components, the culture chamber system includes a culture chamber 10, a flow loop 20, a support base 30, a pump 40, and an electrical actuator 50. The particular arrangement of the components is representative in FIG. 1 , and the components can be arranged in other ways as compared to that shown. The following paragraphs provide a description of the structural and functional aspects of the components shown in FIG. 1 .

[0047] The culture chamber 10 includes a block 1 1 , which forms a type of vessel of the culture chamber 10. In one example, the block 1 1 can be formed from a block of polydimethylsiloxane (PDMS), although other suitable materials, such as other types of silicon-based organic polymers, can be used. The culture chamber 10 also includes a well 12 formed within the block 1 1 , a flexible membrane 13 adhered to the bottom of the block 1 1 and enclosing the well 12 on one side, and posts 14 extending within the well 12. Each of the posts 14 is adhered at one end to the flexible membrane 13. The size of the culture chamber 10 and the well 12 can vary depending upon the type of cells being grown and other factors, and an example size of the well 12 can range from between about 1 .5 to 20 cm, although other sizes are within the scope of the embodiments.

[0048] The well 12 can be formed in any suitable way by cutting or punching a central portion out from within the block 1 1 . Particularly, the well 12 can be formed by cutting or punching completely through from one side of the block 1 1 to an opposite side of the block 1 1 . To enclose the well 12 on the opposite (e.g. , bottom) side of the block 1 1 , the flexible membrane 13, which is preferably a relatively thin membrane of PDMS material, is adhered to the opposite side of the block 1 1 , enclosing the well 12 on one side. The posts 14 are formed separately and each adhered at one end to the flexible membrane 13, to stand upright within the well 12, using PDMS in liquid or gel form, which cures to a relatively rigid, but stable form. When the well 12 is empty (i.e., empty of the media 22), the posts 14 stand upright (or relatively upright) within the well 12.

[0049] The flow loop 20 includes a media reservoir 21 to hold a volume of media 22, an outlet tube 23 for the media 22 to flow from the media reservoir 21 to the culture chamber 10, and an inlet tube 24 for the media 22 to flow from the culture chamber 10 back to the media reservoir 21 . The flow loop 20 also includes a first one-way valve 25 in-line with the outlet tube 23, a second one-way valve 26 in-line with the inlet tube 24, and a variable flow constrictor 27 in-line with the inlet tube 24. As described in further detail below, the first one way value 25 and the second one-way valve 26 permit flow of the media 22 in one direction from the media reservoir 21 , through the outlet tube 23, into the well 12 of the culture chamber 10, and out of the inlet tube 24 back to the media reservoir 21 . The media 22 flows between the media reservoir 21 and the well 12 of the culture chamber 10 due to a combination of gravity and pumping action performed by the pneumatic pump 40 as described below.

[0050] As described in further detail below with reference to FIGS. 3A and 3B, when the media 22 flows from the media reservoir 21 to the well 12, the flexible membrane 13 stretches or distends downward due to the weight of the media 22 in the well 12. The support base 30, which is placed beneath the well 12, buttresses the distention of the flexible membrane 13, preventing distention beyond a certain extent and to ensure tensile stress on the cultured cells within the tissue. When the flexible membrane 13 stretches or distends downward due to the weight of the media 22 in the well 12, the posts 14 also bend outward from each other along with the flexible membrane 13, which conforms to the shape of the support base 30. Thus, when the well 12 is full of the media 22 and the flexible membrane 13 stretches or distends downward to the support base 30, the posts 14 bend outward from each within the well 12.

[0051] The pump 40 includes a pumping device capable of pumping a volume of air or other suitable fluid between the flexible membrane 13 and the support base 30. Particularly, when the pump 40 pumps the air or other fluid between the flexible membrane 13 and the support base 30, the fluid pushes the flexible membrane 13 back up, off the support base 30, against the weight of the media 22 in the well 12. This pumping forces at least a portion of the media 22 out of the well 12, under a pre-defined pressure based on the variable flow constrictor 27, through the inlet tube 24, and back into the media reservoir 21. When the flexible membrane 13 is pushed back up off the support base 30, the posts 14 also bend back toward each other to stand relatively upright. [0052] Thus, as described herein, the culture chamber system is designed to pump the media 22 into and out of the well 12 of the culture chamber 10, periodically, simulating an environment in which the cardiac tissue cells typically grow. The pump 40 stimulates a pumping action in the culture chamber 10 as cardiac tissue cells grow within the well 12 over time, simulating an environment in which the cardiac tissue cells typically grow. The pumping action of the pump 40 also results in the exchange or flow of the media 22 between the media reservoir 21 and the well 12 of the culture chamber 10 using the flow loop 20.

[0053] The electrical actuator 50 includes a device capable of providing an electrical stimulus within the well 12 through one or more electrodes 51 placed within the well 12. The electrodes 51 can be placed at any suitable locations within the well 12 in various embodiments, and the illustration in FIG. 1 is representative. In one case, an electrode 51 can be placed, respectively, at opposite, far sides of each of the posts 14. In other cases, an electrode 51 can be placed, respectively, on a side of each of the posts 14. As described below, the pump 40 and the electrical actuator 50 can be configured to coordinate pumping and electrical stimulus (at the onset of systole), to simulate the environment in which the cardiac tissue cells typically grow.

[0054] FIGs. 21A-21 C show an embodiment of a culture well and a system as described herein with integrated electrodes for the delivery of electrical stimulation.

[0055] FIGS. 2A and 2B further illustrate an example of the culture chamber 10 according to various embodiments described herein. Particularly, FIG. 2A is a representative, top-down photograph of the well 12 within the block 1 1 , and FIG. 2B is a representative, top- down photograph of cardiac tissue 16 formed between the posts 14 within the well 12.

[0056] Referring first to FIG. 2A, the well 12 can be formed in any suitable way by cutting or punching a central portion out from within the block 1 1 . The well 12 can be formed by cutting or punching completely through from one side of the block 1 1 to an opposite side of the block 1 1. The well 12 is shown as an elliptical well, although other shapes can be cut out from within the block 1 1 . To enclose the well 12 on the opposite side of the block 1 1 , the flexible membrane 13 is adhered to the opposite side of the block 1 1 , enclosing the well 12, as described above. The posts 14 are formed separately and each adhered at one end to the flexible membrane 13, to stand upright within the well 12, using PDMS in liquid or gel form, which cures to a relatively rigid, but stable form.

[0057] After the well 12 is formed in the block 1 1 and the flexible membrane 13 is adhered to the back side of the block 1 1 , an agarose 15 or other suitable polymer is filled into the well 12. The agarose 15 gels, molds, or hardens within the well 12 and around the posts 14. As shown in FIG. 2A, a portion of the agarose 15 is then removed from around and between the posts 14, forming a sub-well 12A. A mixture of cardiac cells to be cultured, suspended within a hydrogel, is then filled into the sub-well 12A. The hydrogel can be comprised of any material(s) that support the survival of the cardiac cells, including but not limited to fibrin, collagen, Matrigel, alginate, or a combination of different hydrogels. The cardiac cells within the hydrogel grow into the cardiac tissue 16 formed between the posts 14 within the well 12. Once the cardiac cells grow into the cardiac tissue 16, the remainder of the agarose 15 is then removed from within the well 12 as shown in FIG. 2B. From the point shown in FIG. 2B (or before), the culture chamber system can be operated to simulate the environment in which the cardiac tissue cells typically grow.

[0058] FIG. 3 illustrates a diastolic phase of a pumping process performed using the culture chamber system illustrated in FIG. 1 according to various embodiments described herein. As shown in FIG. 3, the media 22 flows from the media reservoir 21 , through the outlet tube 23 and the one-way valve 25, to the well 12, to simulate a diastolic phase of the pumping process. In this phase, the height or volume of the media 22 within the media reservoir 21 can be adjusted to set the diastolic pressure experienced by the cardiac tissue 16 within the well 1 1.

[0059] During the diastolic phase, the flexible membrane 13 stretches or distends downward due to the weight of the media 22 in the well 12. The support base 30, which is placed beneath the well 12, buttresses the distention of the flexible membrane 13, preventing distention close to the center of the membrane but allowing distention at the edges as shown in FIG. 3. When the flexible membrane 13 stretches or distends downward against the support base due to the weight of the media 22 in the well 12, the posts 14 also bend outward from each other as shown. Thus, the posts 14 bend outward from each within the well 12 during the diastolic phase, stretching the cardiac tissue 16. The pump 40 can be set to generate a vacuum to enhance the stretching or turned off and the electrical actuator 50 can also be turned off during this diastolic phase.

[0060] The media 22 supplies nutrients and other signaling molecules to the cardiac tissue 16 so that the cardiac cells of the cardiac tissue 16 can continue to culture and grow and respond to signaling molecules. The media 22 does not flow from the well 12 to the media reservoir 21 during the diastolic phase, because gravity combined with the resistance provided by the variable constrictor prevents the media 22 from flowing up through the inlet tube 24. Fluid will exit the chamber only if the pressure within the chamber exceeds the combined effects of gravity and the resistance.

[0061] FIG. 4 illustrates a systolic phase of a pumping process performed using the culture chamber system illustrated in FIG. 1 according to various embodiments described herein. During the systolic phase, the pump 40 pumps air or other fluid between the flexible membrane 13 and the support base 30. The fluid pushes the flexible membrane 13 back up (i.e. , contrast with FIG. 3), off the support base 30, against the weight of the media 22 in the well 12 resulting in generation of pressure within the chamber. When the flexible membrane 13 is pushed back up off the support base 30, the posts 14 also bend back toward each other to stand relatively upright as shown in FIG. 4, permitting the cardiac tissue 16 to return to its original relaxed state. The electrical actuator 50 can also supply an electrical pulse to the electrodes 51 at a suitable point in time during the systolic phase to mimic the impulses that drive contraction of the cardiac tissue 16.

[0062] During the systolic phase, the media 22 cannot flow back through the outlet tube 23 because the one-way value 25 does not permit flow of the media 22 back through the outlet tube 23. The pressure experienced within the well 12 during the systolic phase can be set based on adjustment of the variable flow constrictor 27, which can be embodied as an adjustable clamp or other means to constrict a volume of the media 22 that flows through the inlet tube 24.

[0063] As outlined above, the cardiac tissues developed using the culture chamber system can be used for various research and development purposes. Various conditions, including embryonic, adult, diseased, interventions, exercise, etc. can be imposed on cells within the culture chamber system by adjusting the height of the media 22 in the media reservoir 21 , the level or amount of stretch of the flexible membrane 13, the applied pressure from the pump 40, and use of additional pumps or devices to load/unload the culture chamber 10 (e.g., MCS/LVAD devices).

[0064] The culture chamber system can achieve a number of advantages over existing methods, because it is the only platform where the pressure-volume changes seen in any chamber of the heart at any stage of development can be accurately recreated. It offers a combination of electrical and mechanical loading in a single platform and imposes these stresses in a physiologically relevant manner. It can accurately recreate hemodynamics associated with development and disease. It is also the only platform where interventions like mechanical circulatory support (MCS)/LVAD devices can be evaluated. [0065] The features of the embodiments described herein are representative and, in alternative embodiments, certain features and elements can be added or omitted. Additionally, modifications to aspects of the embodiments described herein can be made by those skilled in the art without departing from the spirit and scope of the present invention defined in the following claims, the scope of which are to be accorded the broadest interpretation so as to encompass modifications and equivalent structures.

EXAMPLES

[0066] Now having described the embodiments of the present disclosure, in general, the following Examples describe some additional embodiments of the present disclosure. While embodiments of the present disclosure are described in connection with the following examples and the corresponding text and figures, there is no intent to limit embodiments of the present disclosure to this description. On the contrary, the intent is to cover all alternatives, modifications, and equivalents included within the spirit and scope of embodiments of the present disclosure. The following examples are put forth so as to provide those of ordinary skill in the art with a complete disclosure and description of how to perform the methods and use the probes disclosed and claimed herein. Efforts have been made to ensure accuracy with respect to numbers (e.g., amounts, temperature, etc.), but some errors and deviations should be accounted for. Unless indicated otherwise, parts are parts by weight, temperature is in °C, and pressure is at or near atmospheric. Standard temperature and pressure are defined as 20 °C and 1 atmosphere.

[0067] Example 1

[0068] Abstract

[0069] Human induced pluripotent stem cell (hiPSC)-derived cardio myocytes (hiPSC-CMs) hold great promise for cardiovascular disease modeling and regenerative medicine. However, these cells are both structurally and functionally immature, primarily due to their differentiation into cardiomyocytes occurring under static culture which only reproduces biomolecular cues and ignores the dynamic hemodynamic cues that shape early and late heart development during cardiogenesis. To evaluate the effects of hemodynamic stimuli on hiPSC-CM maturation, the biomimetic cardiac tissue model (BCTM) was utilized to reproduce the hemodynamics and pressure/volume changes associated with heart development. Following 7 days of gradually increasing stimulation, it is shown that hemodynamic loading results in (a) enhanced alignment of the cells and extracellular matrix, (b) significant increases in genes associated with physiological hypertrophy, (c) noticeable changes in sarcomeric organization and potential changes to cellular metabolism, and (d) a significant increase in fractional shortening, suggestive of a positive force frequency response. These findings suggest that culture of hiPSC-CMs under conditions that accurately reproduce hemodynamic cues results in structural organization and molecular signaling consistent with organ growth and functional maturation.

[0070] Introduction

[0071] Cardiovascular disease is the number one cause of death in the USA and throughout the world, accounting for about 31 % of all mortalities [Mozaffarian et al., 2016; Benjamin et al. 2018] The high mortality rate associated with cardiovascular disease can be attributed to the limited intrinsic regenerative capacity of cardio myocytes that make up the heart. Damage and death of cardiomyocytes following events like myocardial infarctions result in the initiation of an inflammatory cascade of events that leads to cardiac tissue remodeling via cardiac fibroblast differentiation into myofibroblasts, and the infarcted areas being replaced with nonfunctional fibrotic scar tissue [Turner and Porter, 2013; Ma et al., 2014] The resulting scar tissue does not contribute to contractile function and results in long term complications that can result in sudden cardiac death, dilated cardiomyopathy, arrhythmias, and heart failure [Pauschinger et al., 1999; Khan and Sheppard, 2006; Dobaczewski et al., 201 1 ; Louzao-Martinez et al., 2016]

[0072] In vitro studies play an important role in modeling diseases and developing drugs and therapeutics. The lack of appropriate human cell types to construct in vitro models of cardiac tissue has limited the relevancy of years of research into mechanisms and potential therapies. A vast majority of these studies used neonatal or embryonic cardiomyocytes from animal and avian sources [Martinsen, 2005; Warkman and Krieg, 2007; Miura and Yelon, 201 1 ] The use of adult cardiomyocytes is limited to acute studies due to the inability to maintain them in an in vitro culture without dedifferentiation [Benardeau et al., 1997; Zhang et al., 2010] The discovery of human induced pluripotent stem cells (hiPSCs) and the advent of protocols to efficiently differentiate them into cardiomyocytes (hiPSC-CMs) represents a major breakthrough in cardiovascular research that, for the first time, provided a human source of cardiomyocytes without the ethical concerns that accompany the use of embryonic stem cells. However, the relevance of hiPSC-CMs is limited, as they represent an immature state that more closely resembles fetal or embryonic cardiomyocytes that are structurally and functionally different from human adult cardiomyocytes [Beilin et al., 2012; Iglesias-Garcia et al., 2013; Nakamura et al. , 2013; Martins et al., 2014; Matsa et al. , 2014; Feric and Radisic, 2016]

[0073] Cardiogenesis and cardiomyocyte maturation involve a highly orchestrated series of events that take place over an extended period of time and continue even after birth. The heart begins to beat and pump blood early in embryogenesis when it is still a linear tube with no formed valves or chambers [Clark and Hu, 1990; Keller et al., 1990] This corresponds with cardiac differentiation of mesodermal precursors into immature cardiomyocytes. These immature cardiomyocytes first develop spontaneous action potentials, and then gradually begin contractions in a random and disorganized fashion sufficient to induce peristaltic blood flow within the linear heart tube. Gradually, these contractions become stronger and more organized, resulting in small changes in pressure and stretch as blood flows through the linear heart tube. The various hemodynamic parameters like ventricular blood pressure, stroke volume, and heart rate increase in proportion to the weight of the embryo, such that blood flow indexed for the weight of the embryo remains constant [Clark et al., 1989; Hu and Clark, 1989] During cardiogenesis, the heart is extremely sensitive to biomechanical cues, such that it will not develop properly in the absence of blood flow, and any alterations to normal hemodynamic stresses result in cardiac malformations and congenital heart defects [Rychter, 1978; Clark et al., 1984; Sedmera et al., 1999; Tobita and Keller, 2000] Therefore, during cardiogenesis, blood flow and associated hemodynamic stresses play a critical role in the formation of the heart and the transformation of naive cardiomyocytes into fully functional adult cardiomyocytes [Srivastava, 2006]

[0074] The dynamic environment and developmental cues associated with blood flow are contrary to the static environment in which hiPSC-CMs are typically differentiated which is devoid of any hemodynamic cues. Moreover, a majority of established protocols achieve differentiation in 2-dimensional (2D) monolayers. The 2D differentiation process under static conditions results in cells that have similar resting membrane potentials, sarcomeric proteins, ion channels, and mechanisms for calcium handling and storage to those in adult cardiomyocytes, but the cells are also smaller in size, exhibit underdeveloped and relatively disorganized sarcomeres, lack t-tubules, possess large pacemaker currents, and beat spontaneously [Itzhaki et al., 2006; O’Hara et al., 201 1 ; Poon et al., 201 1 ; Mummery et al., 2012; Synnergren et al., 2012; Robertson et al., 2013] Using a 2D, biochemically driven process fails to replicate the complexity of the environment in vivo that relies on a combination of extracellular matrix (ECM) cues, electrical cues, and mechanical cues, and also the presence of and interaction with other noncardiomyocyte cell types (including endothelial cells and cardiac fibroblasts) required for card io myocyte maturation [Hirota et al., 1983; Hu and Clark, 1989; Clark and Hu, 1990; Hove et al., 2003; Moorman et al., 2003; Christoffels and Moorman, 2009; Bowers and Baudino, 2010] Therefore, it is conceivable that culturing hiPSC-CMs while reproducing the hemodynamics of cardiogenesis may be the most efficient method of generating mature cardiomyocytes from these PSCs. Several groups have focused on mimicking geometric constraints [Pilarczyk et al., 2016], patterning [Carson et al., 2016], culture in dense collagen matrices [Roberts et al., 2016], culture in 3D fibrin gels [Zhang et al., 2013], and ontomimetic differentiation [Kerscher et al., 2016] to enable the progressive maturation of hiPSCs into hiPSC-CMs. While these efforts resulted in improvements in both the structural and functional maturation of hiPSC-CMs, levels of functional maturation associated with adult cardiomyocytes was not achieved. A recent work by Ronaldson-Bouchard et al. [2018] used progressively increased electromechanical conditioning to mature early-stage hiPSC-CMs and showed superior structural, functional, and metabolic maturation in comparison to other available methods. While this is an impressive achievement and moves the field closer towards an adult cardio myocyte, several functional measures including conduction velocities and contractile force generation fall short of what is seen in the adult myocardium. This could possibly be due to the fact that this process does not recreate the hemodynamic loading associated with gradual increases in pressure/volume loading seen as the heart transitions from a linear heart tube to a neonatal and then a postnatal heart.

[0075] To address the absence of hemodynamics in current protocols, the biomimetic cardiac tissue model (BCTM) was utilized, which is the only system available that can recreate hemodynamic loading and unloading associated with pressure/volume changes in any chamber of the heart at any stage of development. Reproducing the blood flow associated with the early-stage heart was a focus of the work of the present example. A point of novelty of this work is the ability to superimpose the flow, pressure, and stretch associated with critical events during early-stage heart development on iPSC-CMs during in vitro differentiation in a manner where the application of these stresses closely mimics what occurs in vivo. To further enhance physiological significance, early-stage hiPSC-CMs (immediately following mesodermal induction and cardiac differentiation) were co-cultured with neonatal fibroblasts in a 3D fibrin gel to more closely mimic cardiac tissue.

[0076] Materials and Methods

[0077] Cell Culture Chamber Fabrication [0078] Cell culture chambers were fabricated using standard soft lithography using (poly)dimethyl siloxane (PDMS) (Sylgard™ 184, Dow® Corning, Midland, Ml, USA). The chamber was created using a two-step process of bonding a thin PDMS membrane to the bottom of a square PDMS block with an ellipse-shaped hole punched through the center measuring 15 ^c 30 mm. The PDMS block and thin membrane used a ratio of base-to- crosslinker of 10:1 . The thin membrane was formed by spinning 10: 1 PDMS to approximately 80 pm on a silicon wafer within a tabletop centrifuge retro-fitted to hold and center the silicon wafers. Both components were cured overnight at 60 ° C. Once cured, both pieces were cleaned and bonded using oxygen plasma (Harrick Plasma Systems, Ithaca, NY, USA), optimized to create a permanent bond using 700 mTorr pressure, 30 W, and 20 s of exposure. Next, two 10:1 PDMS posts measuring 6.5 mm in height and 2 mm in diameter were placed on top of the thin membrane and along the ellipse’s long axis 12 mm apart. The posts were bonded by placing the bonded PDMS chambers on top of a hotplate set to 150 ° C, and they were then dipped in liquid 10: 1 PDMS and placed through a removable guide that allowed for consistent placement. After autoclaving, the chambers were filled with 2% hot liquid agarose and allowed to cool and solidify. Circles around each post and a center lane connecting these circles were punched out, leaving a dumb-bell shape within the chamber (FIG. 5A). Finally, the chamber was ready for seeding with fibrin-encapsulated cells. The final product, once the gel was added and the agarose removed can be seen in FIG. 5B.

[0079] Human iPSC-CM Culture and Differentiation

[0080] Human iPSCs were cultured within 6-well plates coated for 1 h at room temperature with human embryonic stem-cell (hESC)-qualified Matrigel® supplied by Corning (354277) at the manufacturer-recommended concentration for maintaining pluripotency. The cells were cultured using MTeSR™1 (Stemcell Technologies™ 85850) and receive daily media changes. At 75-85% confluency, the cells were disassociated using Accutase™ (Stemcell Technologies 07920), and then reseeded in MTeSR™1 supplemented with 10 mM Y-27632, ROCK1 inhibitor (Selleckchem S1049) at a split ratio of 1 : 6. After 12 h, hiPSCs were switched to MTeSR™1 without ROCK1 inhibitor. After 3 days, hiPSCs reached a confluency of 75-85% and were switched to Roswell Park Memorial Institute (RPMI) medium (Gibco™ 1 1875093) supplemented with B27 without insulin (Gibco™ A18956-01) and 10 pM WNT promoter CHIR99021 (Stemcell Technologies™ 72052), considered day 1 of differentiation. After exactly 24 h, the cells were washed and switched to media without CHIR99021 . They were maintained with daily media changes until day 4, when the media was supplemented with 5 mM WNT inhibitor IWP-2 (Stemcell Technologies™ 72124). They were cultured with IWP-2 until day 6. On day 8, the cells were switched to RPMI medium supplemented with -B27 with insulin (Gibco™ 17504044). At this point, the cells were referred to as hiPSC-CMs and had begun spontaneously beating. Medium was changed daily until day 10 and refreshed thereafter every 2 days. Only plates that contained > 60% beating hiPSC-CMs were used for further study.

[0081] Human Primary Fibroblast Culture

[0082] Neonatal primary dermal fibroblasts were purchased from American Type Culture Collection (ATCC®) (PCS-201-010) and cultured in primary fibroblast medium (ATCC® PCS-201-030) supplemented with Fibroblast Growth Kit- Low serum (ATCC® PCS- 201-041 ). Cells were cultured according to ATCC® recommendations and were not used in any study past passage 15.

[0083] Fibrin Gel Encapsulation

[0084] On day 12, the hiPSC-CMs were disassociated by using 0.25% Trypsin/EDTA and combined with human fibroblasts disassociated by using 0.05% Trypsin/EDTA at a final density of 2.0 ´ 106 fibroblasts and 4.0 ´ 106 hiPSC-CMs per 300 mL of fibrin solution. The fibrin solution was composed of Dulbecco’s Modified Eagle’s Medium-High Glucose (DMEM- HG) supplemented with 10% fetal bovine serum (FBS), 50 pg/mL L-ascorbic acid (Sigma® A5960), 1 × nonessential amino acids (Gibco™ 1 1 140-050), and 5 mg/mL aminocaproic acid (Acros Organics™ 103301000), and contained a final concentration of 2 mg/mL fibrinogen (Millipore® 341576), 0.9 units thrombin (Millipore® 605195)/mg of fibrin (0.54 units/gel), and 2 mM CaCI ₂ . The fibrin-encapsulated cells were then placed within a 5% CO ₂ incubator for 30 min at 37°C to complete the fibrin polymerization. Following polymerization and until the end of experimentation, the cells were maintained with DMEM-HG supplemented with 10% FBS, 50 pg/mL ascorbic acid, 1 ´ nonessential amino acids, and 5 mg/mL aminocaproic acid. Once within the fibrin gel, the cells were maintained by daily media changes. Cells were allowed to spread and adapt to the 3D architecture as well as recover from the trypsinization and encapsulation process for 5 days before beginning BCTM stimulation.

[0085] BCTM Stimulation

[0086] The mechanism of action of the BCTM revolves around using dynamic pressure differences above and below the cell culture chamber membrane to reproduce the cardiac cycle [Rogers et al.,2016]. The BCTM comprises a media reservoir set to a specified height to create a downward filling pressure imposed onto the cell culture member to mimic a diastolic filling pressure. A programmable pneumatic pump can be used to increase the pressure below the cell culture membrane to generate a systolic pressure. The cycling of the pneumatic pump coupled with one-way flow control values allows the actuation of the culture membrane to displace fluid in one direction with each contraction (FIG. 5E). The BCTM has changed from its previous published iterations through its adaptation to the culture of 3D tissue fibers suspended between 2 posts (FIGs. 5A, 5B). Another notable difference in the current design is the presence of a dome-shaped insert below the cell culture chamber and aligned between the 2 posts, that deforms the membrane as it is pushed down onto the insert during the diastole phase of the cycle. By recording the post movements during the cardiac cycle (FIG. 5C, 5D), it is possible to calculate strain.

[0087] BCTM Stimulation Conditions

[0088] To adapt to the mechanical stress, the cells are placed under gentle conditions and ramped up over time (FIG. 6). Gradually increasing the stress was found to be necessary for hiPSC-CMs to survive under hemodynamic loading [Rogers et al., 2016]. On day 0, the stimulation was set at a peak systolic pressure of 5 mm Hg, an end-diastolic pressure of 1 mm Hg, and a maximum strain of 0.2%. These were increased to 10 mm Hg, 3 mm Hg, and 0.7%, respectively, after 24 h (day 1 ); 15 mm Hg, 5 mm Hg, and 1.2%, respectively, on day 2; and 30 mm Hg, 10 mm Hg, and 2.3%, respectively, on day 3. This final condition was maintained until the end of the 7-day experiment. Throughout the entire experiment, the cycle frequency was maintained at 1 Hz so as to replicate a normal human resting heart rate. The pressure and stretch stimulus were determined to ensure cell survival; higher levels of mechanical stimulation have been found to be detrimental to cell survival. The 7-day stimulation represented the initial phase of embryonic heart development during which blood flow and pressure/volume changes are initiated. Also, 7 days is sufficient time to clearly see changes in cardiomyocyte structural and functional maturation.

[0089] Staining

[0090] At the end of the experiment, the fibrin-encapsulated cells were removed from the BCTM, and then fixed using fresh 4% paraformaldehyde at room temperature for 1 h before moving to 4°C. Samples were then placed in Optimal Cutting Temperature (OCT) solution and frozen for cryosectioning of 10-mm-thick slices. Sections were immunolabeled with rabbit anti-cardiac troponin T (anti-cTnT; Abeam®, ab91605) and mouse monoclonal anti-a sarcomeric actin (anti-a-SA) antibody (Sigma®, A2172). Nuclei were counterstained with DAPI. Images were taken using an Olympus IX83 epifluorescence microscope. [0091] Transmission Electron Microscopy

[0092] At the end of the 7-day stimulation protocol, the cell fibers were removed from the BCTM, and then fixed using a fixative composed of 2% paraformaldehyde and 2.5% glutaraldehyde in a 0.1 -M sodium cacodylate solution for 1 h at room temperature, and then moved to 4°C overnight. Samples were then taken to the UAB high-resolution imaging facility where they were further processed, sliced, embedded, and sectioned. Slices were imaged using a Tecnai™ Spirit T12 transmission electron microscope.

[0093] RNA qRT-PCR

[0094] After 7 days of stimulation, the cells were removed from the BCTM, carefully slid off the posts, placed within a liquid nitrogen safe cryovial with 700 mL Trizol (Invitrogen), and then flash-frozen and maintained in liquid nitrogen until processing. Once enough samples had been collected for gene analysis, they were thawed at room temperature and immediately placed within a beadmill (BeadBug™ D1030) containing 1 .0 mm silica beads (Benchmark D1031-10) and milled for 30 s at speed 4,000. RNA was then isolated using standard phenol-chloroform extraction methods for quantification of the transcripts GADPH, GATA4, NKX2.5, MYLK, MYL2, KCNH2, CAMK2B, and MYH7. Complementary DNA for mRNAs was obtained from 2 mg total RNA in a 20-pL reaction using TaqMan™ reverse- transcription reagents (Thermo Fisher Scientific) and 100 pmol of oligo(dT)15 primer. This mixture was incubated at 37°C for 2 h. Quantitative RT-PCR was performed with primers (Table 1 ) using the Eppendorf RealPlex2 real-time PCR system. Complementary DNA synthesized from 100 ng total RNA was combined with Maxima SYBR Green qPCR Master Mix (Thermo Fisher Scientific) and 0.5 pM each of the forward and reverse primers. Cycling conditions were as follows: 95°C for 10 min followed by 40 cycles of amplification (95°C denaturation for 15 s, and 60°C annealing-extension for 1 min). To avoid the influence of genomic contamination, the forward and reverse primers for each gene were located in

Table 1 : Primers for qRT-PCR Gene Expression Analysis

[0095] different exons. Relative quantitative expression of genes of interest was calculated after normalizing to the housekeeping gene glyceraldehyde 3-phosphate dehydrogenase ( GAPDH ).

[0096] Fractional Shortening

[0097] Immediately after removing the hiPSC-CM fibers from the BCTM, their intrinsic (spontaneous) and paced contractions at 1.0, 1 .3, 1 .5, and 2.0 Hz were recorded using a video camera. A Myopacer field stimulator (lonoptix) with adapted carbon electrodes provided field stimulation at 4.0 V/cm with a 40-ms duration. The field was set to alternate polarity to avoid build-up on the electrodes. The carbon electrodes were placed near the bottom of the cell culture surface and between the post and outside edge of the cell culture chamber so that the field ran along the entire length of the fiber. HBSS (Gibco 14025-092) warmed to 37°C was used during the recordings.

[0098] Fractional shortening was estimated using video analysis and Lagrangian strain calculations for finitely small displacement, as previously described by Bonet and Wood [2008] and used by our group [Nguyen et al. , 2009, 2015] Briefly, if p1 and p2 are 2 nearby points defined on the first frame and d1 denotes the Euclidian distance between p1 and p2 at the first frame (i.e. , reference frame), then dj denotes the Euclidian distance between these 2 points at time frame j. The engineering strain S, at frame j, is then defined as:

[0099] S _j = (d _j — d _i )/ d ₁ (eq. 1 )

[0100] A custom algorithm was written to estimate planar Lagrangian strains. First, the peak displacement was determined by tracking 2 points based on their appearance features through the image frames, and then calculating the displacement between the 2 points over the image frames. The maximum displacement was then selected as the reference frame. Given that the contractions are unidirectional due to alignment of the fibers, 2 pairs of points on the first (reference) frame were determined and their tracked pairs on the selected frame with maximum displacement. Lagrangian strain for each of the points was estimated using the previously mentioned equation. Finally, the mean Lagrangian strain was calculated and expressed as % shortening. To account for drift due to macroscopic movement of the fiber, the analysis was performed on a per-beat basis where the measurements correspond to maximum and minimum displacements per contraction. [0101] Statistical Analysis

[0102] An unpaired t test was used for the analysis of the gene expression and fractional shortening data with two-tailed significance set at p < 0.05 for both experiments.

[0103] Results

[0104] Cardiomyocyte Sarcomeric Structure

[0105] Immunofluorescence microscopy was performed to visualize sarcomeric organization in both stimulated and static samples. Images show that both stimulated and static samples had organized sarcomeres with striated organization of the a-SA and cTnT (FIGs. 7A-7B).

[0106] Cardiomyocyte Ultrastructural Morphology

[0107] Transmission electron microscopy (TEM) was performed to evaluate changes in intracellular and extracellular structure and organization. Images clearly show that the ECM associated with the stimulated samples is more aligned and composed of longer, thicker fibrils (FIG. 8B, top) in comparison to the shorter, thinner, and more disorganized fibrils in static controls (FIG. 8A, top). Static control cells were inundated with lipid deposits and unorganized regions of sarcomeric proteins (denoted with a black and white arrow, respectively; FIG. 8A, middle, left) whereas the stimulated samples did not show lipid deposits but had more highly organized sarcomeres (FIG. 8B, middle). A highly magnified image of sarcomeres shows that Z-disks, A-bands, and l-bands are clearly visible in both the static controls and stimulated samples (FIGs. 8A-8B, bottom) with evidence of the beginnings of the formation of an M-line in the stimulated sample (black arrow).

[0108] Expression of Genes Involved in Structural and Functional Maturation

[0109] Static and stimulated samples were profiled for changes in genes associated with cardiomyocyte hypertrophy and cardiac-specific structural maturation (FIGs. 9A-9G). Several genes associated with hypertrophic growth were upregulated in the stimulated samples. Results confirm that GATA4, myosin regulatory light chain 2 ( MYL2 ), calcium/calmodulin-dependent protein kinase type II b chain ( CAMK2B ), and myosin heavy chain b ( MYH7) were significantly upregulated in stimulated samples versus in the static controls. No statistically significant changes were found in NKX2.5, myosin light chain kinase (, MYLK ), and potassium voltage-gated channel subfamily H member 2 ( KCNH2 ). Although a mixed population of fibroblasts and cardio myocytes make up the cell samples used for gene analysis, the genes chosen were highly cardiac specific. [0110] Functional Maturation of Cardiomyocytes

[0111] Video analysis of the fractional shortening shows that stimulated samples achieved greater fractional shortening than the samples cultured under static conditions. It can also be noted that the static samples demonstrated a decrease in fractional shortening at progressively higher pacing frequencies, while that of the stimulated samples peaked at the 1 .3-Hz pacing frequency before progressively decreasing at higher pacing frequencies (FIG. 10). Importantly, significantly higher fractional shortening was measured in the stimulated samples (n = 6) in comparison to the static samples (n = 5) at the frequencies 1.3, 1 .5, and 2.0 Hz.

[0112] Discussion/Conclusion

[0113] Heart development is a complicated multistage process that requires careful coordination of genetic, biochemical, physical and electrical cues to enable transformation of a monolayer of cells into a complex multi-chambered organ capable of pumping blood to the entire body. Given that the primary role of the heart is to pump blood, it is essential that the developing heart not only generates contractile forces but can also adapt to mechanical stresses associated with hemodynamic loading and unloading. Mechanical stresses are experienced by cardiomyocytes during embryonic heart development when blood flow is first established following development of the linear heart tube. This corresponds with cardiac differentiation of mesodermal precursors into immature cardiomyocytes. These immature cardiomyocytes first develop spontaneous action potentials and gradually begin contractions in a random and disorganized fashion which is sufficient to induce peristaltic blood flow within the linear heart tube. Gradually these contractions become stronger and more organized resulting in small changes in pressure and stretch as blood flows through the linear heart tube.

[0114] Following looping, chamber formation is initiated with the growth and establishment of valves and synchronous and organized contractions begin to appear following establishment of the sinoatrial and arterioventricular nodes. Various hemodynamic parameters like ventricular blood pressure, stroke volume and heart rate are tightly regulated.51 Cardiac output increases in proportion to the weight of the embryo such that blood flow indexed for the weight of the embryo remains constant23 and growth is accelerated in response to experimentally increased developed pressure.24 During development, the heart is extremely sensitive to biomechanical cues and constant interaction between hemodynamic stresses developed due to cardiac tissue motion and blood flow, shape and direct heart formation.52 The heart does not develop properly in the absence of blood flow39,52 and any alterations to normal hemodynamic stresses have been shown to result in cardiac malformations and congenital heart defects.25-28

[0115] Given the importance of mechanical loading during embryonic heart development, it was sought to determine if hemodynamic stimulation could potentially impact the differentiation of hiPSC-CMs into a more mature phenotype. This involved the culture of early-stage (immediately following cardiac differentiation of mesodermal precursors) hiPSC- CMs under peristaltic flow and pressure, stretch which is gradually increased to 20 mm Hg, and 3% stretch within the BCTM. The results provide preliminary validation that recreating mechanical loading during early cardiac differentiation induces changes associated with the stage of embryonic heart development linked to an increase in chamber size and functional maturation. Immunofluorescence microscopy provided clear evidence of physiological hypertrophy and enhanced organization of hiPSC-CMs subject to stimulation within the BCTM (FIGs. 7A-7B). Visualization of cTnT and a-SA showed significant improvements in the sarcomeric alignment and localization of actin and troponin in clear, alternating patterns in the stimulated samples, but not in the static controls. It also appears that each sarcomere was wider and longer than the sarcomeres in the static controls, although this was not quantified.

[0116] Examination of ultrastructural morphology using TEM revealed important distinctions between the static controls and stimulated samples (FIG. 8A-8B). It appears that mechanical stimulation greatly enhanced the alignment of the ECM and enabled the formation of longer and thicker ECM fibrils. This is potentially beneficial to cardiomyocyte maturation by providing alignment cues and coordinating force transmission along the length of the engineered muscle fiber. Stimulation also increased the overall alignment of sarcomeres and prevented disorganized regions that were observed in the static controls. High-magnification images showed organization of sarcomeres including Z-disks, A-bands, and l-bands in both static controls and stimulated samples. Closer examination showed the beginnings of the formation of M-lines in the stimulated samples. Surprisingly, lipid deposits were found within hiPSC-CMs in the static controls. The absence of lipid deposits in the stimulated samples could possibly be explained by a switch to fatty acid metabolism due to the higher energetic demand of stimulated cells to achieve physiological hypertrophy and growth. However, at this stage of development (embryonic/neonatal), glycolysis is still the primary source of energy associated with proliferation and organ growth. Additional studies are necessary to determine the significance of lipid accumulation in the static controls and evaluate possible differences in metabolic activity between the static controls and stimulated samples. TEM is commonly used to evaluate ultrastructure morphology of cardiomyocytes. Adult cardiomyocytes are defined by organized bands and lines (A-band, M-line, Z-line, and l-bands) that make up the sarcomere and are flanked by sarcoplasmic reticulum and a high number of mitochondria that typically run along the length of the sarcomere. The presence of each band along with increased alignment places stimulated samples closer to what would be expected from fetal cardiac tissue [Ronaldson-Bouchard et al. , 2018], which demonstrating that there is early and progressive maturation.

[0117] Gene expression studies provide further evidence of physiological hypertrophy in the BCTM-stimulated samples (FIGs. 9A-9G). The upregulation of GATA4 is significant due to its presence in developing and adult cardiomyocytes where it has been linked to hypertrophic growth of both immature and mature cardiomyocytes. MYL2 and MYH7 are both part of the myosin complex involved in binding to actin filaments and are integral to proper contractile function. Upregulation of these 2 genes implies the creation of new sarcomeric structures and possibly explains the increase in fractional shortening observed in the stimulated samples. Human CAMK2B is a calcium handling protein that was upregulated in the stimulated samples, and this suggests improved calcium handling and hypertrophy. The lack of a change in KCNH2 could be due to the stimulation period being too short or that electrical stimulation may also be required for upregulation of voltage-regulated channels, as seen in other studies using electrical stimulation for maturation [Eng et al., 2016]

[0118] Finally, the fractional shortening measurements suggested significant improvements in percent shortening of stimulated samples in comparison to static controls (FIG. 10). Interestingly, the fractional shortening data suggests that stimulation within the BCTM induced a positive force frequency response in the stimulated samples, i.e., an increase in contractile force with increasing frequency. The stimulated samples demonstrated a peak fractional shortening at 1 .3 Hz which gradually decreased with increasing frequencies, creating a bell curve indicative of the force frequency response measured in mature cardiac tissue. A similar bell curve peaking at 1.5 Hz is present in normal healthy cardiac tissue. It is possible that this is an example of a positive force frequency response, but it should be noted that the increases in fractional shortening between pacing frequencies in the stimulated samples were not significantly different. However, it is clear that the fractional shortening was significantly increased in the stimulated samples when compared to the static culture at most pacing frequencies. [0119] In summary, the BCTM was used to successfully recreate hemodynamic loading associated with early heart development. 3D-engineered cardiac tissue fibers constructed using early-stage hiPSC-CMs/neonatal fibroblast cocultures stimulated for 7 days within the BCTM showed noticeable and quantifiable changes in structure, alignment, organization, hypertrophic gene expression, and function when compared to engineered fibers maintained under static culture. These results suggest that the BCTM can be used to recreate hemodynamic cues associated with embryonic development and could possibly be used in conjunction with electrical stimulation to mimic conditions associated with the postnatal (adult) heart.

[0120] References (Alphabeticall

[0121] 2017. "Cardiovascular diseases (CVDs)." World Health Organization, Last

Modified 5/17/2017, accessed 6/5/2018. http://www.who.int/en/news-room/fact- sheets/detail/cardiovascular-diseases-(cvds).

[0122] Beilin, Milena, Maria C Marchetto, Fred H Gage, and Christine L Mummery. 2012. "Induced pluripotent stem cells: the new patient?" Nature reviews Molecular cell biology 13 (1 1 ):713.

[0123] Benardeau, Agnes, Stephane N. Hatem, Catherine Riicker-Martin, Sophie Tessier, Sylvie Dinanian, Jane-Lyse Samuel, Edouard Coraboeuf, and Jean-Jaques Mercadier. 1997. "Primary Culture of Human Atrial Myocytes is Associated with the Appearance of Structural and Functional Characteristics of Immature Myocardium." Journal of Molecular and Cellular Cardiology 29 (5): 1307-1320. doi: https://doi.Org/10.1006/jmcc.1996.0366.

[0124] Benjamin, Emelia J., Salim S. Virani, Clifton W. Callaway, Alexander R. Chang, Susan Cheng, Stephanie E. Chiuve, Mary Cushman, Francesca N. Delling, Rajat Deo, Sarah D. de Ferranti, Jane F. Ferguson, Myriam Fornage, Cathleen Gillespie, Carmen R. Isasi, Monik C. Jimenez, Lori Chaffin Jordan, Suzanne E. Judd, Daniel Lackland, Judith H. Lichtman, Lynda Lisabeth, Simin Liu, Chris T. Longenecker, Pamela L. Lutsey, David B. Matchar, Kunihiro Matsushita, Michael E. Mussolino, Khurram Nasir, Martin O’Flaherty, Latha P. Palaniappan, Dilip K. Pandey, Mathew J. Reeves, Matthew D. Ritchey, Carlos J. Rodriguez, Gregory A. Roth, Wayne D. Rosamond, Uchechukwu K.A. Sampson, Gary M. Satou, Svati H. Shah, Nicole L. Spartano, David L. Tirschwell, Connie W. Tsao, Jenifer H. Voeks, Joshua Z. Willey, John T. Wilkins, Jason HY. Wu, Heather M. Alger, Sally S. Wong, and Paul Muntner. 2018. "Heart Disease and Stroke Statistics— 2018 Update: A Report From the American Heart Association." Circulation doi: 10.1 161/cir.0000000000000558.

[0125] Bonet, Javier, and Richard D Wood. 1997. Nonlinear continuum mechanics for finite element analysis: Cambridge university press.

[0126] Bowers, Stephanie LK, and Troy A Baudino. 2010. "Laying the groundwork for growth: cell-cell and cell-ECM interactions in cardiovascular development." Birth Defects Research Part C: Embryo Today: Reviews 90 (1 ): 1 -7.

[0127] Carson, Daniel, Marketa Hnilova, Xiulan Yang, Cameron L. Nemeth, Jonathan H. Tsui, Alec S. T. Smith, Alex Jiao, Michael Regnier, Charles E. Murry, Candan Tamerler, and Deok-Ho Kim. 2016. "Nanotopography-Induced Structural Anisotropy and Sarcomere Development in Human Cardiomyocytes Derived from Induced Pluripotent Stem Cells." ACS Applied Materials & Interfaces doi: 10.1021/acsami.5b1 1671.

[0128] Christoffels, Vincent M., and Antoon F.M. Moorman. 2009. "Development of the Cardiac Conduction System: Why Are Some Regions of the Heart More Arrhythmogenic Than Others?" Circulation: Arrhythmia and Electrophysiology 2 (2): 195-207. doi: 10.1 161/circep.108.829341 .

[0129] Claire, Robertson, Tran David D., and George Steven C. 2013. "Concise Review: Maturation Phases of Human Pluripotent Stem Cell-Derived Cardiomyocytes." STEM CELLS 31 (5):829-837. doi: doi: 10.1002/stem.1331 .

[0130] Clark, E. B., and N. Hu. 1990. "Hemodynamics of the developing cardiovascular system." Ann N Y Acad Sci 588:41 -7.

[0131] Clark, E. B., N. Hu, P. Frommelt, G. K. Vandekieft, J. L. Dummett, and R. J. Tomanek. 1989. "Effect of increased pressure on ventricular growth in stage 21 chick embryos." Am J Physiol 257 (1 Pt 2):H55-61.

[0132] Clark, E. B., N. Hu, and G. C. Rosenquist. 1984. "Effect of conotruncal constriction on aortic-mitral valve continuity in the stage 18, 21 and 24 chick embryo." Am J Cardiol 53 (2):324-7.

[0133] Dobaczewski, M., W. Chen, and N. G. Frangogiannis. 201 1. "Transforming growth factor (TGF)-beta signaling in cardiac remodeling." J Mol Cell Cardiol 51 (4):600-6. doi: 10.1016/j.yjmcc.2010.10.033.

[0134] Feric, Nicole T, and Milica Radisic. 2016. "Maturing human pluripotent stem cell-derived cardiomyocytes in human engineered cardiac tissues." Advanced drug delivery reviews 96:1 10-134. [0135] Hirota, A., T. Sakai, S. Fujii, and K. Kamino. 1983. "Initial development of conduction pattern of spontaneous action potential in early embryonic precontractile chick heart." Dev Biol 99 (2):517-23.

[0136] Hove, J. R., R. W. Koster, A. S. Forouhar, G. Acevedo-Bolton, S. E. Fraser, and M. Gharib. 2003. "Intracardiac fluid forces are an essential epigenetic factor for embryonic cardiogenesis." Nature 421 (6919):172-7. doi: 10.1038/nature01282.

[0137] Hu, N., and E. B. Clark. 1989. "Hemodynamics of the stage 12 to stage 29 chick embryo." Circ Res 65 (6): 1665-70.

[0138] Iglesias-Garcia, O., B. Pelacho, and F. Prosper. 2013. "Induced pluripotent stem cells as a new strategy for cardiac regeneration and disease modeling." J Mol Cell Cardiol 62:43-50. doi: 10.1016/j.yjmcc.2013.04.022.

[0139] ILANIT, ITZHAKI, SCHILLER JACKIE, BEYAR RAFAEL, SATIN JONATHAN, And GEPSTEIN LIOR. 2006. "Calcium Handling in Embryonic Stem Cell-Derived Cardiac Myocytes." Annals of the New York Academy of Sciences 1080 (1 ):207-215. doi: doi:10.1 196/annals.1380.017.

[0140] Keller, B. B., N. Hu, and E. B. Clark. 1990. "Correlation of ventricular area, perimeter, and conotruncal diameter with ventricular mass and function in the chick embryo from stages 12 to 24." Circ Res 66 (1 ): 109-14.

[0141] Kerscher, Petra, Irene C. Turnbull, Alexander J. Hodge, Joonyul Kim, Dror Seliktar, Christopher J. Easley, Kevin D. Costa, and Elizabeth A. Lipke. 2016. "Direct hydrogel encapsulation of pluripotent stem cells enables ontomimetic differentiation and growth of engineered human heart tissues." Biomaterials 83:383-395. doi: http://dx.doi.Org/10.1016/j.biomaterials.2015.12.01 1 .

[0142] Khan, R., and R. Sheppard. 2006. "Fibrosis in heart disease: understanding the role of transforming growth factor-beta in cardiomyopathy, valvular disease and arrhythmia." Immunology 1 18 (1): 10-24. doi: 10.1 1 1 1/j.1365-2567.2006.02336.x.

[0143] Louzao-Martinez, Laura, Aryan Vink, Magdalena Harakalova, Folkert W. Asselbergs, Marianne C. Verhaar, and Caroline Cheng. 2016. "Characteristic adaptations of the extracellular matrix in dilated cardiomyopathy." International Journal of Cardiology 220:634-646. doi: https://doi.Org/10.1016/j.ijcard.2016.06.253.

[0144] Ma, Yonggang, Lisandra E. de Castro Bras, Hiroe Toba, Rugmani Padmanabhan Iyer, Michael E. Hall, Michael D. Winniford, Richard A. Lange, Suresh C. Tyagi, and Merry L. Lindsey. 2014. "Myofibroblasts and the extracellular matrix network in post-myocardial infarction cardiac remodeling." Pfiugers Archiv - European Journal of Physiology 466 (6):1 1 13-1 127. doi: 10.1007/s00424-014-1463-9.

[0145] Martins, A. M., G. Vunjak-Novakovic, and R. L. Reis. 2014. "The Current Status of iPS Cells in Cardiac Research and Their Potential for Tissue Engineering and Regenerative Medicine." Stem Cell Rev 10 (2): 177-90. doi: 10.1007/s12015-013-9487-7.

[0146] Martinsen, B. J. 2005. "Reference guide to the stages of chick heart embryology." Dev Dyn 233 (4):1217-37. doi: 10.1002/dvdy.20468.

[0147] Matsa, Elena, Paul W Burridge, and Joseph C Wu. 2014. "Human stem cells for modeling heart disease and for drug discovery." Science translational medicine 6 (239):239ps6-239ps6.

[0148] Miura, G. I., and D. Yelon. 201 1 . "A guide to analysis of cardiac phenotypes in the zebrafish embryo." Methods Cell Biol 101 :161 -80. doi: 10.1016/b978-0-12-387036- 0.00007-4.

[0149] Moorman, Antoon, Sandra Webb, Nigel A. Brown, Wouter Lamers, and Robert H. Anderson. 2003. "DEVELOPMENT OF THE HEART: (1 ) FORMATION OF THE CARDIAC CHAMBERS AND ARTERIAL TRUNKS." Heart 89 (7):806-814.

[0150] Mozaffarian, D., E. J. Benjamin, A. S. Go, D. K. Arnett, M. J. Blaha, M. Cushman, S. R. Das, S. de Ferranti, J. P. Despres, H. J. Fullerton, V. J. Howard, M. D. Huffman, C. R. Isasi, M. C. Jimenez, S. E. Judd, B. M. Kissela, J. H. Lichtman, L. D. Lisabeth, S. Liu, R. H. Mackey, D. J. Magid, D. K. McGuire, E. R. Mohler, 3rd, C. S. Moy, P. Muntner, M. E. Mussolino, K. Nasir, R. W. Neumar, G. Nichol, L. Palaniappan, D. K. Pandey, M. J. Reeves, C. J. Rodriguez, W. Rosamond, P. D. Sorlie, J. Stein, A. Towfighi, T. N. Turan, S. S. Virani, D. Woo, R. W. Yeh, and M. B. Turner. 2016. "Heart Disease and Stroke Statistics- 2016 Update: A Report From the American Heart Association." Circulation 133 (4):e38-360. doi: 10.1 161/cir.0000000000000350.

[0151] Mummery, C. L., J. Zhang, E. S. Ng, D. A. Elliott, A. G. Elefanty, and T. J. Kamp. 2012. "Differentiation of human embryonic stem cells and induced pluripotent stem cells to cardio myocytes: a methods overview." Circ Res 1 1 1 (3):344-58. doi: 10.1 161/circresaha.1 10.227512.

[0152] Nakamura, K., K. Hirano, and S. M. Wu. 2013. "iPS cell modeling of cardiometabolic diseases." J Cardiovasc Transl Res 6 (1 ):46-53. doi: 10.1007/s12265-012- 9413-4. [0153] Nguyen, M. D., G. Giridharan, S. D. Prabhu, and P. Sethu. 2009. "Microfluidic cardiac circulation model (microCCM) for functional cardiomyocyte studies." ConfProc IEEE Eng Med Biol Soc 2009:1060-3. doi: 10.1 109/iembs.2009.5334997.

[0154] Nguyen, M. D., J. P. Tinney, F. Ye, A. A. Elnakib, F. Yuan, A. El-Baz, P. Sethu, B. B. Keller, and G. A. Giridharan. 2015. "Effects of Physiologic Mechanical Stimulation on Embryonic Chick Cardiomyocytes Using a Microfluidic Cardiac Cell Culture Model." Anal Chem. doi: 10.1021/ac503716z.

[0155] O'Hara, Thomas, Laszlo Virag, Andras Varro, and Yoram Rudy. 201 1. "Simulation of the Undiseased Human Cardiac Ventricular Action Potential: Model Formulation and Experimental Validation." PLOS Computational Biology 7 (5):e1002061 . doi: 10.1371/journal. pcbi.1002061.

[0156] Pauschinger, M., D. Knopf, S. Petschauer, A. Doerner, W. Poller, P. L. Schwimmbeck, U. Kuhl, and H. P. Schultheiss. 1999. "Dilated cardiomyopathy is associated with significant changes in collagen type I/Ill ratio." Circulation 99 (21):2750-6.

[0157] Pilarczyk, Gotz, Alexandra Raulf, Manuel Gunkel, Bernd Fleischmann, Robert Lemor, and Michael Hausmann. 2016. "Tissue-Mimicking Geometrical Constraints Stimulate Tissue-Like Constitution and Activity of Mouse Neonatal and Human-Induced Pluripotent Stem Cell-Derived Cardiac Myocytes." Journal of Functional Biomaterials 7 (1 ):1 .

[0158] Poon, Ellen, Chi-wing Kong, and Ronald A. Li. 201 1 . "Human Pluripotent Stem Cell-Based Approaches for Myocardial Repair: From the Electrophysiological Perspective." Molecular Pharmaceutics 8 (5): 1495-1504. doi: 10.1021/mp2002363.

[0159] Roberts, Meredith A., Dominic Tran, Kareen L. K. Coulombe, Maria Razumova, Michael Regnier, Charles E. Murry, and Ying Zheng. 2016. "Stromal Cells in Dense Collagen Promote Cardiomyocyte and Microvascular Patterning in Engineered Human Heart Tissue." Tissue Engineering Part A 22 (7-8):633-644. doi: 10.1089/ten. tea.2015.0482.

[0160] Rogers, Aaron J, Vladimir G Fast, and Palaniappan Sethu. 2016. "Biomimetic Cardiac Tissue Model Enables the Adaption of Human Induced Pluripotent Stem Cell Cardiomyocytes to Physiological Hemodynamic Loads." Analytical Chemistry 88 (19):9862- 9868.

[0161] Ronaldson-Bouchard, Kacey, Stephen P Ma, Keith Yeager, Timothy Chen, LouJin Song, Dario Sirabella, Kumi Morikawa, Diogo Teles, Masayuki Yazawa, and Gordana Vunjak-Novakovic. 2018. "Advanced maturation of human cardiac tissue grown from pluripotent stem cells." Nature 556 (7700):239.

[0162] Rychter, Z. 1978. "Analysis of relations between aortic arches and aorticopulmonary septation." Birth Defects Orig Artie Ser 14 (7):443-8.

[0163] Sedmera, D., T. Pexieder, V. Rychterova, N. Hu, and E. B. Clark. 1999. "Remodeling of chick embryonic ventricular myoarchitecture under experimentally changed loading conditions." Anat Rec 254 (2):238-52.

[0164] Srivastava, D. 2006. "Genetic regulation of cardiogenesis and congenital heart disease." Annu Rev Pathol 1 : 199-213. doi: 10.1 146/annurev.pathol.1.1 10304.100039.

[0165] Synnergren, Jane, Caroline Ameen, Andreas Jansson, and Peter Sartipy.

2012. "Global transcriptional profiling reveals similarities and differences between human stem cell-derived cardiomyocyte clusters and heart tissue." Physiological Genomics 44 (4):245-258. doi: 10.1 152/physiolgenomics.001 18.201 1.

[0166] Tobita, K., and B. B. Keller. 2000. "Right and left ventricular wall deformation patterns in normal and left heart hypoplasia chick embryos." Am J Physiol Heart Circ Physiol 279 (3):H959-69.

[0167] Turner, Neil A, and Karen E Porter. 2013. "Function and fate of myofibroblasts after myocardial infarction." Fibrogenesis & tissue repair 6 (1 ):5.

[0168] Warkman, A. S., and P. A. Krieg. 2007. "Xenopus as a model system for vertebrate heart development." Semin Cell Dev Biol 18 (1 ):46-53. doi:

10.1016/j.semcdb.2006.1 1 .010.

[0169] Zhang, D., I. Y. Shadrin, J. Lam, H. Q. Xian, H. R. Snodgrass, and N. Bursae.

2013. "Tissue-engineered cardiac patch for advanced functional maturation of human ESC- derived cardiomyocytes." Biomaterials 34 (23):5813-20. doi:

10.1016/j.biomaterials.2013.04.026.

[0170] Zhang, Yiqiang, Tao-Sheng Li, Shuo-Tsan Lee, Kolja A Wawrowsky, Ke Cheng, Giselle Galang, Konstantinos Malliaras, M Roselle Abraham, Charles Wang, and Eduardo Marban. 2010. "Dedifferentiation and proliferation of mammalian cardiomyocytes." P/oS one 5 (9):e12559.

[0171] Example 2

[0172] Abstract [0173] Objective: Cardiovascular research and regenerative strategies have been significantly limited by the lack of relevant cell culture models that can recreate complex hemodynamic stresses associated with pressure-volume changes in the heart. Methods: To address this issue, a Biomimetic Cardiac Tissue Chip (CTC) Model was designed where encapsulated cardiac cells can be cultured in 3D fibres and subjected to hemodynamic loading to mimic pressure-volume changes seen in the left ventricle. These 3D fibres are suspended within a microfluidic chamber between two posts and integrated within a flow loop. Various parameters associated with heart function like heart rate, peak-systolic pressure, end-diastolic pressure and volume, end-systolic pressure and volume, and duration ratio between systolic and diastolic can all be precisely manipulated allowing culture of cardiac cells under developmental, normal, and disease states. Results: The CTC can significantly impact cardiovascular research by reproducing the pathophysiological mechanical stresses associated with pressure overload and volume overload as shown in the example below. Results using H9c2 cells, a cardiomyogenic cell line, clearly show that culture within the CTC under pathological hemodynamic loads accurately induces morphological and gene expression changes similar to that seen in both hypertrophic and dilated cardiomyopathy. Under pressure overload the cells within the CTC see increased hypertrophic remodelling and fibrosis whereas cells subject to prolonged volume overload experience significant changes to cellular aspect ratio through thinning and elongation of the engineered tissue. Conclusions: These results demonstrate that the CTC can be used to create highly relevant models where hemodynamic loading and unloading are accurately reproduced for cardiovascular disease modelling.

[0174] I. Introduction

[0175] Cardiac Tissue Chips (CTCs) can potentially be used as models of the heart to understand signaling mechanisms involved in cardiac development and disease and as models for cardiovascular drug development and testing. The heart constantly interacts with blood and any drug delivered systemically interacts with the myocardium. Given the importance of heart (pump) function, drug-induced cardiotoxicity (arrhythmia risk and compromised contractile function) has been a major reason for pharmaceutical withdrawal of FDA approved drugs [1] Historically, the development of Cardiac Tissue Chip Models has been limited due to lack of in vitro cell culture systems that can reproduce hemodynamic loading and unloading associated with heart (pump) function. Hemodynamic loads are extremely important in embryonic heart development and in progression of cardiovascular disease and play a major role in both physiological hypertrophy (organ growth and maturation) as well as pathological hypertrophy (adverse remodeling and compromised myocardial function).

[0176] Cardiac cells have been cultured primarily using standard cell culture techniques in static conditions. While these techniques can provide biochemical stimuli, they fail to adequately replicate critical hemodynamic (mechanical) stresses experienced by cardiac cells in vivo. Commercially available systems include the FlexCell® technology which imposes stretch stimulus on cultured cells, and lonOptix systems that provide the ability to either electrically stimulate or combine electrical stimulation with stretch. There have also been several groups that have developed technologies to mimic various aspects of the in vivo environment to enable differentiation and maturation of stem cell derived cardiomyocytes including mimicking geometric constraints [2], nanopatterning [3], culture in dense collagen matrices [4], culture in 3D fibrin gels [5], use of thyroid hormone and miRNAs [6, 7], use of passive or dynamic stretch and electrical stimulation [8], moderate afterload [9], ontomimetic differentiation [10] and electromechanical stimulation [1 1 ] While these approaches have yielded varying levels of success from a standpoint of maturation and differentiation, these approaches still do not mimic the complex pressure-volume (PV) changes observed in the left-ventricle of the heart which is the location associated with most common cardiac pathologies. To address this issue the cardiac cell culture model (CCCM) was previously developed [12, 13] and the biomimetic cardiac tissue model (BCTM) [14] platforms which mimic PV loading associated with any chamber of the heart at any stage of development. However, these approaches failed to mimic the predominantly uniaxial stretch mechanics and three dimensional (3D) and multicellular architecture associated with the ventricular wall. To develop the next generation of cardiac tissue culture systems, the Cardiac Tissue Chip (CTC) was developed which can not only reproduce uniaxial loading of and left ventricular PV loading associated with normal and pathological cardiac function but also accomplish culture of multi-cellular 3D tissue. Within the CTC, cardiomyocytes and fibroblasts are organized as 3D tissue within fibrin gels suspended between two posts. The engineered 3D tissue experiences cycles of stretch, contraction (pressure), ejection and relaxation similar to that observed during the cardiac cycle.

[0177] In cardiovascular disease, progression of pathological remodeling occurs in large part due to sustained non-physiological hemodynamic loads. The two most common manifestations of non-physiological hemodynamic loads include pressure and volume overload. Pressure overload is associated with an increase in left-ventricular pressure that typically develops due to chronic hypertension as a consequence elevated systemic resistance and leads to hypertrophic cardiomyopathy (HCM) [15] HCM manifests as uncontrolled hypertrophic growth of myocytes typically in an unorganized manner that leads to concentric cardiomyocyte hypertrophy (thickening) along with apoptosis [16-18] Resulting cell death and hypertrophy coincide with increased fibrosis mediated by cardiac fibroblasts, oxidative stress and structural remodeling [19-21 ] The increased matrix deposition disrupts normal cardiac conduction leading to increases in arrhythmia risks, and increased wall stiffness and thickness [22] Cardiac fibroblasts play a central role in pathological hypertrophy seen in pressure overload by resisting the increases in wall stress through activation of TGF- b signaling and increased extracellular matrix (ECM) deposition [23] Left ventricular volume overload occurs due to diastolic dysfunction or mitral regurgitation resulting in excess residual volume in the left ventricle causing thinning of the ventricular wall which leads to dilated cardiomyopathy (DCM) and ultimately heart failure [24] Thinning of the ventricle leads to elongation of cardiomyocytes in the left ventricular wall that undergo eccentric hypertrophy. [25] The increase in stretch has been linked to matrix degradation, increase in reactive oxygen species (ROS), and disruption of cytoskeletal filament proteins involved in mechanotransduction. [26-29]

[0178] Given the importance of hemodynamic stresses as a consequence of PV changes within the left-ventricle, it is essential that disease models of cardiac tissue have the ability to accurately recreate pathological hemodynamic loads. While several systems and models exist to subject cultured cardiac cells to pressure or stretch [12, 13, 30, 31], none of these models mimic the complexity of loading associated with the cardiac cycle. The CTC therefore address the absence of relevant in vitro systems to recreate cardiac physiology and pathophysiology resulting from hemodynamic imbalances. By recreating the hemodynamics associated with the left-ventricle in vivo during both normal heart function and during cardiomyopathies (FIG. 1 1), it may be possible to develop new and relevant models of cardiovascular disease. Using these models, the mechanotransduction pathways that result in pathological remodeling seen in hypertrophic and dilated cardiomyopathy can be targeted for discovery and testing of new therapeutics. In this example the use of the CTC to create models of both pure pressure and pure volume overload is detailed, and it is shown that critical structural and functional changes associated with both these conditions can be reproduced. For model development the h9c2 rat myoblast cell line was chosen which is widely used for cardiovascular disease modelling and can be maintained in culture for prolonged duration. The choice of h9c2 cells is relevant for modelling pressure overload as prior studies confirm that these cells are indeed responsive to hypertrophic signaling in vitro and appropriate to study the effects of pathological levels of pressure [32-34] Despite the fact that these cells do not contract either spontaneously or under the influence of external pacing, these cells possess similar calcium channels, sarcomeric proteins, and metabolic profiles to cardiac cells [35-37]. H9c2 cells have also been used extensively to study myocardial infarctions and reperfusion injuries due to their similarities in comparison to in vivo responses to oxidative stress [38, 39] In the context of this study, these cells provide an ideal cell type to accomplish development of cardiovascular disease models.

[0179] II. Materials and Methods

[0180] Cell Culture Chamber Fabrication: Cell culture chambers were fabricated using standard soft-lithography using (poly)dimethyl siloxane (PDMS) (Sylgard™ 184, Dow Corning, Midland, Ml) using standard techniques in the laboratory [40] Assembled chambers were autoclaved and filled with 2% molten agarose and allowed to cool and solidify. A dumbbell shape pattern was created via removal of agarose using a hole punch and scalpel. Finally, the chamberwas seeded with fibrin encapsulated cells and allowed to cross-link. The agarose mold was removed following 24 hours in culture resulting in a cell laden fibrin gel suspended between two posts. Following hemodynamic stimulation the cell laden fibrin gel can be removed from the post and retains its form making further analysis simpler (FIG. 12A).

[0181] CTC Stimulation: The setup and working cycle of the CTC are the same as described previously. [41 ] Briefly, dynamic pressure differences above and below the cell culture chamber membrane in conjunction with directional flow control valves were exploited to reproduce the cardiac cycle. This version is different from the prior model as it was adapted for culture of 3D tissue fibers suspended between two posts (FIGs. 12B-12C).

[0182] Cell Culture and Fibrin Gel Encapsulation: 2.2x10 ⁶ h9c2 and 1 .0x10 ⁶ primary rat cardiac fibroblasts (RFB) cells were mixed together into 300ul_ media containing a final concentration of 2mg/ml_ fibrinogen (Millipore® 341576), 0.9 units thrombin (Millipore® 605195) per mg fibrin (0.54 units/gel), and 2mM CaCI ₂ . The gels were allowed to polymerize for 30 minutes within the incubator before media was added. During the gel making process, and until completion of the experiment, the culture medium was supplemented with 5mg/ml_ aminocaproic acid (Acros Organics® 103301000) to inhibit fibrinolysis.

[0183] RNA qRT-PCR: After 48 hours of stimulation the cells are removed from the CTC and placed within 700mL Trizol and then flash frozen and maintained in liquid nitrogen until processing. Once all samples were collected, they were thawed at room temperature and immediately placed within a beadmill (BeadBug D1030) containing 1 .0mm silica beads (Benchmark D1031 -10) and milled for 30 seconds at speed 4000. RNA is then isolated using standard phenol-chloroform extraction methods, and cDNA is created using Maxima kit (K1671 ). Custom ThermoFisher Taq-Man® array plates (4391528) were created under design ID RAAAADM and coupled with Applied Biosystems master mix (4369514). Plates were run on a QuantStudio 3.

[0184] Statistical Analysis

[0185] An unpaired t-test was used for the analysis of the gene expression and aspect ratio data with two-tailed significance set at p < 0.05.

[0186] III. Results

[0187] Pressure-Volume Changes Seen in Normal and Pathological Conditions:

Using the CTC pressure-volume changes associated with normal, pure pressure overload and pure volume overload could be accurately replicated. Normal conditions were set to a 1 Hz cardiac cycle with l OOmmHg peak-systolic pressure, 10mmHg end-diastolic pressure, and a minimum strain of 0% and maximum strain of 2% (FIG. 13A). Pressure overload conditions were set to a 1 Hz cardiac cycle with 160mmHg peak-systolic pressure, 10mmHg end-diastolic pressure, and a minimum strain of 0% and a maximum strain of 2% (FIG. 13B). Volume overload conditions were set to a 1 Hz cardiac cycle with l OOmmHg peak-systolic pressure, 30mmHg end-diastolic pressure, and minimum strain of 2% and maximum strain of 7% (FIG. 13C). To reproduce the hemodynamics of dilated cardiomyopathy the volume overload was not allowed to return to its resting level of strain in order to reproduce the constant wall stress seen in vivo.

[0188] Gross Tissue Morphology and Organization: At the conclusion of the experiment, fibers were evaluated for alignment and general tissue morphology using H&E staining of paraffin embedded slices of the fibers (FIGs. 14A-14D). In all samples, regardless of the presence or absence of stimulation, the fiber contraction around the post caused alignment of the ECM resulting in cell alignment. H&E staining confirmed that cells were able to spread and evenly distribute throughout the tissue fiber. To determine if the experimental conditions were able to induce an increase in ECM deposition the slices of were also stained using Masson’s trichrome stain (FIGs. 15A-15H).

[0189] It is clear both from the 20X and 40X images that 48-hour stimulation within the BCTM under conditions of pressure overload results in a noticeable increase in collagen deposition along the centre of the fibre. [0190] Tissue Aspect Ratio after 7 Days of Stimulation: While changes in gene expression manifest within 48 hours, it was necessary to extend the duration of stimulation to 7 days to observe noticeable changes in tissue dimensions. Following 7 days of stimulation, the volume overload samples experienced a significant increase (18%) in total length of the fibre and this change can be seen in the kink in the fibre when returned to the original post position (FIG. 16). Measurement of the width of the fibres was used in conjunction with overall fibre length to estimate aspect ratios of the tissue fibres for each condition. The change in aspect ratios for the pressure overload and volume overload were significant when compared to the static control and normal sample (FIG. 16). These results show a striking similarity to in vivo observations where pressure overload induces an increase in aspect ratio (width/length) and volume overload induces thinning and elongation, thus decreasing the aspect ratio. For clarity, the longest dimension is considered the length in these measurements.

[0191] Gene Expression Profiling: Following the 48 hours of CTC stimulation, the changes in gene expression were profiled for static (N = 3), normal (N =3), pressure overload (N = 3) and volume overload (N =3). Results show that there are indeed significant changes in the expression of various genes following culture under normal and pathological loading (FIGs. 17A-17N). Genes were classified into four categories within the array. Profiling of genes associated with fibrosis found that Collagen I was significantly upregulated in the pressure overload samples whereas transforming growth factor beta (TGF-b) was significantly upregulated in both pressure overload and normally stimulated samples. Evaluation of genes associated with cytoskeletal filament proteins found that desmin was significantly upregulated in the volume overload samples. Finally, evaluation of genes associated with matrix remodelling and antioxidant signalling found that tissue inhibitors of matrix metalloproteases (TIMPs) 1 and 2 were upregulated under pressure overload conditions and a significant increase in superoxide dismutase (SOD) 1 and SOD2 in both pressure and volume overload samples. Data is represented as relative change with the expression of each gene normalized to GAPDH expression levels.

[0192] IV. Discussion

[0193] Cardiovascular disease modelling, drug testing and drug discovery have been significantly impacted by the absence of relevant in-vitro cell culture models to culture cardiac cells under physiological and pathophysiological pressure-volume loading. Cardiomyocytes and other cardiac cells rely on these mechanical stress signals to maintain homeostasis and ensure normal heart function. However, chronic exposure to pathological levels of either pressure or stress can result in activation of compensatory mechanisms that lead to maladaptive cardiac tissue remodeling. To accurately model cardiomyopathies associated with altered mechanical loading as in the cases of pure pressure or volume overload, it is essential to evaluate cardiac cells in an environment where pressure-volume changes seen in the dysfunctional myocardium can be accurately recreated. The CTC is a unique model system that accurately reproduces mechanical loads that drive progression of cardiac tissue remodeling and was used in this study to establish conditions of both pressure and volume overload to determine if maladaptive remodeling and molecular signaling mechanisms that drive structural and functional changes associated with either of these conditions can be accurately replicated in an in-vitro model system.

[0194] To enable mechanical stimulation of engineered 3D tissue, a device and method for suspending 3D fibrin fibres composed of cardiac cells between a set of rigid polymeric posts mounted onto a thin polymeric membrane was developed. Using this setup cellular alignment through the inherent compaction of the fibrin gel prior to imposition of any cyclic mechanical strain was achieved. This setup also allows for the application of uniaxial strain to the cells along their long axis during CTC stimulation. The natural alignment of the cells within the fibres ensured that the only difference between the stimulated and static samples was the imposition of hemodynamic loads.

[0195] Pressure overload is associated with concentric hypertrophy (thickening of myocytes), increased fibrosis, oxidative stress, and inflammation. After 7 days of stimulation under conditions of pressure overload (160 mmHg, 2% strain) confirms that stimulated tissue undergoes hypertrophy, becoming thicker with an increased aspect ratio. Pressure overload also resulted in a significant increase in collagen deposition within the fibres as seen with Masson trichrome staining. This increase in collagen deposition was reinforced with gene expression studies confirming a statistically significant (~ 7-fold) increase in collagen I gene expression in comparison to static controls. Matrix metalloproteinases (MMPs) are responsible for the degradation of ECM proteins and tissue inhibitors of metalloproteinases (TIMPs) inhibit the ability of MMPs to degrade the ECM. Gene expression results show that both TIMP1 and TIMP2 were significantly upregulated in pressure overload samples whereas a statistically significant increase in MMPs was not observed suggesting that the change in the balance between TIMPs and MMPs favouring TIMPs indicates an environment favouring fibrosis in pressure overload. These results are similar to the pathophysiology of cardiac tissue remodeling in vivo where pressure overload induces a stiffening of the myocardial wall through collagen deposition and a thickening through myocyte hypertrophy. TGF-b is potent pro-fibrotic signaling molecule and is upregulated in both pressure and volume overload in vivo and results in increased ECM deposition and the differentiation of fibroblast into myofibroblasts [42-50] In studies it was found that TGF-b was significantly upregulated only in the pressure overload and normal samples and not in the volume overload samples. The upregulation of TGF-b in the normal samples is most likely due to stress induced pathways; studies have shown that stretching cells can induce increases in TGF-b expression [51 , 52]

[0196] Volume overload is associated with the thinning of the ventricularwall resulting in eccentric hypertrophy (myocyte thinning), significant matrix degradation, oxidative stress, and inflammation. Following 48 hours of stimulation under conditions of volume overload, no statistically significant difference in overall fibre length in comparison to the other conditions was observed. The absence of a change in length of the volume overload sample after 48 hours led us to extend the duration of stimulation to 7 days. Following 7 days of stimulation under conditions of volume overload within the BCTM, the fibres were significantly longer (~ 18%) than the static, normal, and pressure samples. The increased stimulation time had no effect on the final lengths of the static, normal, and pressure samples. The resulting increase in length of the volume overload samples demonstrates the reproduction of the elongated tissue phenotype seen in vivo and suggests that progression of eccentric hypertrophy associated with ventricular wall thinning is a time dependent phenomenon. As expected, desmin was significantly upregulated in the volume overload samples. Desmin is an intermediate filament protein that functions as an important mechanosensor that connects the sarcomere through its Z-disk to the mitochondria, nucleus, ECM, through costameres and intercalated disks, and other organelles [53-55] In vivo studies show that volume overload causes desmin to become damaged and disorganized resulting in increased transcriptional regulation [56, 57] It is possible that volume overload induced damage to the cytoskeleton causes the activation of a compensatory mechanism to replace the damaged/broken down desmin [58, 59]

[0197] Both volume and pressure overload induced upregulation of SOD gene expression. SOD1 is typically localized in the cytoplasm whereas SOD2 in the located in the mitochondria. Hypertrophic and dilated cardiomyopathies have both been linked to an increase in intracellular ROS in vivo [60-64] Several studies have shown that an unchecked increase in intracellular ROS can lead to the tissue and cell damage seen in cardiomyopathies [65-67] It is likely that the observed increase in in SOD at the transcript level is linked to a compensatory antioxidant response in cells subject to pressure and volume overload.

[0198] V. Conclusions

[0199] In conclusion, provided herein is evidence that the CTC can serve as a valuable model system via reproduction of the fundamental characteristics of hemodynamic stress associated with pure volume and pressure overload. These results clearly show that major structural and molecular changes associated with pressure and volume overload can be recreated and provide strong rationale for development of in vitro models of cardiac tissue where the pressure-volume associated with the cardiac cycle are used to study cardiovascular disease pathologies. While this study utilized H9c2 cells and primary cardiac fibroblasts, more relevant cells types including human induced pluripotent stem cell derived cardiomyocytes (hiPSC-CMs) and human embryonic stem cell derived cardio myocytes (hESC-CMs) cultured within the CTC could be used to model human cardiovascular disease and cardiac tissue remodelling in response to changes in hemodynamics.

[0200] References

[0201] [1 ] B. Fermini, J. C. Hancox, N. Abi-Gerges, M. Bridgland-Taylor, K. W.

Chaudhary, T. Colatsky, K. Cornell, W. Crumb, B. Damiano, G. Erdemli, G. Gintant, J. Imredy, J. Koerner, J. Kramer, P. Levesque, Z. Li, A. Lindqvist, C. A. Obejero-Paz, D. Rampe, K. Sawada, D. G. Strauss, and J. I. Vandenberg,“A New Perspective in the Field of Cardiac Safety Testing through the Comprehensive In Vitro Proarrhythmia Assay Paradigm,” J Biomol Screen, vol. 21 , no. 1 , pp. 1 -1 1 , Jan, 2016.

[0202] [2] G. Pilarczyk, A. Raulf, M. Gunkel, B. Fleischmann, R. Lemor, and M.

Hausmann,“Tissue-Mimicking Geometrical Constraints Stimulate Tissue-Like Constitution and Activity of Mouse Neonatal and Human-Induced Pluripotent Stem Cell-Derived Cardiac Myocytes,” Journal of Functional Biomaterials, vol. 7, no. 1 , pp. 1 , 2016.

[0203] [3] D. Carson, M. Hnilova, X. Yang, C. L. Nemeth, J. H. Tsui, A. S. T.

Smith, A. Jiao, M. Regnier, C. E. Murry, C. Tamerler, and D.-H. Kim, “Nanotopography- Induced Structural Anisotropy and Sarcomere Development in Human Cardiomyocytes Derived from Induced Pluripotent Stem Cells,” ACS Applied Materials & Interfaces, 2016/02/1 1 , 2016.

[0204] [4] M. A. Roberts, D. Tran, K. L. K. Coulombe, M. Razumova, M. Regnier,

C. E. Murry, and Y. Zheng,“Stromal Cells in Dense Collagen Promote Cardiomyocyte and Microvascular Patterning in Engineered Human Heart Tissue,” Tissue Engineering Part A, vol. 22, no. 7-8, pp. 633-644, 2016/04/01 , 2016.

[0205] [5] D. Zhang, I. Y. Shadrin, J. Lam, H. Q. Xian, H. R. Snodgrass, and N.

Bursae,“Tissue-engineered cardiac patch for advanced functional maturation of human ESC- derived cardiomyocytes,” Biomaterials, vol. 34, no. 23, pp. 5813-20, Jul, 2013.

[0206] [6] S. S. Parikh, D. J. Blackwell, N. Gomez-Hurtado, M. Frisk, L. Wang, K.

Kim, C. P. Dahl, A. Fiane, T. Tonnessen, D. O. Kryshtal, W. E. Louch, and B. C. Knollmann, “Thyroid and Glucocorticoid Hormones Promote Functional T-Tubule Development in Human-Induced Pluripotent Stem Cell-Derived Cardiomyocytes,” Circ Res, vol. 121 , no. 12, pp. 1323-1330, Dec 8, 2017.

[0207] [7] K. T. Kuppusamy, D. C. Jones, H. Sperber, A. Madan, K. A. Fischer,

M. L. Rodriguez, L. Pabon, W.-Z. Zhu, N. L. Tulloch, X. Yang, N. J. Sniadecki, M. A. Laflamme, W. L. Ruzzo, C. E. Murry, and H. Ruohola-Baker,“Let-7 family of microRNA is required for maturation and adult-like metabolism in stem cell-derived cardiomyocytes,” Proceedings of the National Academy of Sciences, vol. 1 12, no. 21 , pp. E2785-E2794, 2015.

[0208] [8] J. L. Ruan, N. L. Tulloch, M. V. Razumova, M. Saiget, V. Muskheli, L.

Pabon, H. Reinecke, M. Regnier, and C. E. Murry, “Mechanical Stress Conditioning and Electrical Stimulation Promote Contractility and Force Maturation of Induced Pluripotent Stem Cell-Derived Human Cardiac Tissue,” Circulation, vol. 134, no. 20, pp. 1557-1567, Nov 15, 2016.

[0209] [9] A. Leonard, A. Bertero, J. D. Powers, K. M. Beussman, S. Bhandari, M. Regnier, C. E. Murry, and N. J. Sniadecki, “Afterload promotes maturation of human induced pluripotent stem cell derived cardiomyocytes in engineered heart tissues,” Journal of Molecular and Cellular Cardiology, vol. 1 18, pp. 147-158, 2018/05/01/, 2018.

[0210] [10] P. Kerscher, I. C. Turnbull, A. J. Hodge, J. Kim, D. Seliktar, C. J.

Easley, K. D. Costa, and E. A. Lipke,“Direct hydrogel encapsulation of pluripotent stem cells enables ontomimetic differentiation and growth of engineered human heart tissues,” Biomaterials, vol. 83, pp. 383-395, 311, 2016.

[0211] [1 1 ] K. Ronaldson-Bouchard, S. P. Ma, K. Yeager, T. Chen, L. Song, D.

Sirabella, K. Morikawa, D. Teles, M. Yazawa, and G. Vunjak-Novakovic, “Advanced maturation of human cardiac tissue grown from pluripotent stem cells,” Nature, vol. 556, no. 7700, pp. 239-243, Apr, 2018. [0212] [12] M. D. Nguyen, J. P. Tinney, F. Ye, A. A. Elnakib, F. Yuan, A. El-Baz,

P. Sethu, B. B. Keller, and G. A. Giridharan,“Effects of physiologic mechanical stimulation on embryonic chick cardio myocytes using a microfluidic cardiac cell culture model,” Anal Chem, vol. 87, no. 4, pp. 2107-13, Feb 17, 2015.

[0213] [13] M. D. Nguyen, J. P. Tinney, F. Yuan, T. J. Roussel, A. El-Baz, G.

Giridharan, B. B. Keller, and P. Sethu,“Cardiac cell culture model as a left ventricle mimic for cardiac tissue generation,” Anal Chem, vol. 85, no. 18, pp. 8773-9, Sep 17, 2013.

[0214] [14] A. J. Rogers, V. G. Fast, and P. Sethu,“Biomimetic Cardiac Tissue

Model Enables the Adaption of Human Induced Pluripotent Stem Cell Cardiomyocytes to Physiological Hemodynamic Loads,” Anal Chem, vol. 88, no. 19, pp. 9862-9868, Oct 4, 2016.

[0215] [15] N. Matsushita, N. Ishida, M. Ibi, M. Saito, A. Sanbe, H. Shimojo, S.

Suzuki, H. Koepsell, Y. Takeishi, Y. Morino, E. Taira, Y. Sawa, and M. Hirose, “Chronic Pressure Overload Induces Cardiac Hypertrophy and Fibrosis via Increases in SGLT1 and IL-18 Gene Expression in Mice,” Int Heart J, Aug 1 1 , 2018.

[0216] [16] P. Anversa, J. M. Capasso, G. Olivetti, and E. H. Sonnenblick,“Cellular basis of ventricular remodeling in hypertensive cardiomyopathy,” Am J Hypertens, vol. 5, no. 10, pp. 758-70, Oct, 1992.

[0217] [17] B. Schwartzkopff, W. Motz, M. Vogt, and B. E. Strauer,“Heart failure on the basis of hypertension,” Circulation, vol. 87, no. 5 Suppl, pp. lv66-72, May, 1993.

[0218] [18] B. Swynghedauw, “Remodeling of the heart in chronic pressure overload,” Basic Res Cardiol, vol. 86 Suppl 1 , pp. 99-105, 1991.

[0219] [19] E. E. Creemers, and Y. M. Pinto,“Molecular mechanisms that control interstitial fibrosis in the pressure-overloaded heart,” Cardiovasc Res, vol. 89, no. 2, pp. 265- 72, Feb 1 , 201 1 .

[0220] [20] S. K. Maulik, and S. Kumar,“Oxidative stress and cardiac hypertrophy: a review,” Toxicol Mech Methods, vol. 22, no. 5, pp. 359-66, Jun, 2012.

[0221] [21 ] K. T. Weber, R. Pick, J. E. Jalil, J. S. Janicki, and E. P. Carroll,

“Patterns of myocardial fibrosis,” J Mol Cell Cardiol, vol. 21 Suppl 5, pp. 121-31 , Dec, 1989.

[0222] [22] R. Khan, and R. Sheppard,“Fibrosis in heart disease: understanding the role of transforming growth factor-beta in cardiomyopathy, valvular disease and arrhythmia,” Immunology, vol. 1 18, no. 1 , pp. 10-24, May, 2006. [0223] [23] K. Fujiu, and R. Nagai, “Fibroblast-mediated pathways in cardiac hypertrophy,” Journal of Molecular and Cellular Cardiology, vol. 70, pp. 64-73, 2014/05/01/, 2014.

[0224] [24] B. A. Mohamed, M. Schnelle, S. Khadjeh, D. Lbik, M. Herwig, W. A.

Linke, G. Hasenfuss, and K. Toischer,“Molecular and structural transition mechanisms in long-term volume overload,” Eur J Heart Fail, vol. 18, no. 4, pp. 362-71 , Apr, 2016.

[0225] [25] Y. Liu, A. R. Dillon, M. Tillson, C. Makarewich, V. Nguyen, L. Dell'ltalia,

A. K. Sabri, V. Rizzo, and E. J. Tsai,“Volume overload induces differential spatiotemporal regulation of myocardial soluble guanylyl cyclase in eccentric hypertrophy and heart failure,” J Mol Cell Cardiol, vol. 60, pp. 72-83, Jul, 2013.

[0226] [26] J. Barnes, and L. J. Dell'ltalia,“The multiple mechanistic faces of a pure volume overload: implications for therapy,” Am J Med Sci, vol. 348, no. 4, pp. 337-46, Oct, 2014.

[0227] [27] S. Reddy, and D. Bernstein, “Molecular Mechanisms of Right

Ventricular Failure,” Circulation, vol. 132, no. 18, pp. 1734-42, Nov 3, 2015.

[0228] [28] J. Barnes, B. Pat, Y. W. Chen, P. C. Powell, W. E. Bradley, J. Zheng,

A. Karki, X. Cui, J. Guichard, C. C. Wei, J. Collawn, and L. J. Dell'ltalia,“Whole-genome profiling highlights the molecular complexity underlying eccentric cardiac hypertrophy,” Ther Adv Cardiovasc Dis, vol. 8, no. 3, pp. 97-1 18, Mar 31 , 2014.

[0229] [29] J. D. Gladden, M. I. Ahmed, S. H. Litovsky, C. G. Schiros, S. G. Lloyd,

H. Gupta, T. S. Denney, Jr., V. Darley-Usmar, D. C. McGiffin, and L. J. Dell'ltalia,“Oxidative stress and myocardial remodeling in chronic mitral regurgitation,” Am J Med Sci, vol. 342, no. 2, pp. 1 14-9, Aug, 201 1 .

[0230] [30] A. M. De Jong, A. H. Maass, S. U. Oberdorf-Maass, R. A. De Boer, W.

H. Van Gilst, and I. C. Van Gelder, “Cyclical stretch induces structural changes in atrial myocytes,” J Cell Mol Med, vol. 17, no. 6, pp. 743-53, Jun, 2013.

[0231] [31 ] M. Ferroni, S. Giusti, D. Nascimento, A. Silva, F. Boschetti, and A.

Ahluwalia, “Modeling the fluid-dynamics and oxygen consumption in a porous scaffold stimulated by cyclic squeeze pressure,” Med Eng Phys, vol. 38, no. 8, pp. 725-32, Aug, 2016.

[0232] [32] S. J. Watkins, G. M. Borthwick, and H. M. Arthur,“The H9C2 cell line and primary neonatal cardiomyocyte cells show similar hypertrophic responses in vitro,” In Vitro Cellular & Developmental Biology-Animal, vol. 47, no. 2, pp. 125-131 , 201 1 . [0233] [33] C. Vindis, R. D’Angelo, E. Mucher, A. Negre-Salvayre, A. Parini, and

J. Mialet-Perez,“Essential role of TRPC1 channels in cardiomyoblasts hypertrophy mediated by 5-HT 2A serotonin receptors,” Biochemical and biophysical research communications, vol. 391 , no. 1 , pp. 979-983, 2010.

[0234] [34] F. Xie, W. Liu, F. Feng, X. Li, L. Yang, D. Lv, X. Qin, L. Li, and L. Chen,

“A static pressure sensitive receptor APJ promote H9c2 cardiomyocyte hypertrophy via PI3K- autophagy pathway,” Acta Biochim Biophys Sin, vol. 46, no. 8, pp. 699-708, 2014.

[0235] [35] J. Hescheler, R. Meyer, S. Plant, D. Krautwurst, W. Rosenthal, and G.

Schultz,“Morphological, biochemical, and electrophysiological characterization of a clonal cell (H9c2) line from rat heart,” Circulation research, vol. 69, no. 6, pp. 1476-1486, 1991 .

[0236] [36] K. R. Sipido, and E. Marban, “L-type calcium channels, potassium channels, and novel nonspecific cation channels in a clonal muscle cell line derived from embryonic rat ventricle,” Circulation research, vol. 69, no. 6, pp. 1487-1499, 1991 .

[0237] [37] A. F. Branco, S. P. Pereira, S. Gonzalez, O. Gusev, A. A. Rizvanov, and P. J. Oliveira, “Gene expression profiling of H9c2 myoblast differentiation towards a cardiac-like phenotype,” P/oS one, vol. 10, no. 6, pp. e0129303, 2015.

[0238] [38] Z.-X. Shan, Q.-X. Lin, Y.-H. Fu, C.-Y. Deng, Z.-L. Zhou, J.-N. Zhu, X.-

Y. Liu, Y.-Y. Zhang, Y. Li, and S.-G. Lin,“Upregulated expression of miR-1/miR-206 in a rat model of myocardial infarction,” Biochemical and biophysical research communications, vol. 381 , no. 4, pp. 597-601 , 2009.

[0239] [39] X. Wu, L. He, F. Chen, X. He, Y. Cai, G. Zhang, Q. Yi, M. He, and J.

Luo, “Impaired autophagy contributes to adverse cardiac remodeling in acute myocardial infarction,” PLoS One, vol. 9, no. 1 1 , pp. e1 12891 , 2014.

[0240] [40] V. Parichehreh, K. Medepallai, K. Babbarwal, and P. Sethu,

“Microfluidic inertia enhanced phase partitioning for enriching nucleated cell populations in blood,” Lab Chip, vol. 13, no. 5, pp. 892-900, Mar 7, 2013.

[0241] [41 ] A. J. Rogers, V. G. Fast, and P. Sethu,“Biomimetic Cardiac Tissue

Model Enables the Adaption of Human Induced Pluripotent Stem Cell Cardiomyocytes to Physiological Hemodynamic Loads,” Analytical Chemistry, vol. 88, no. 19, pp. 9862-9868, 2016.

[0242] [42] M. Dobaczewski, W. Chen, and N. G. Frangogiannis,“Transforming growth factor (TGF)-beta signaling in cardiac remodeling,” J Mol Cell Cardiol, vol. 51 , no. 4, pp. 600-6, Oct, 201 1 . [0243] [43] J.-M. Li, and G. Brooks, “Differential Protein Expression and

Subcellular Distribution of TGF i , 2and 3in Cardiomyocytes During Pressure Overload- induced Hypertrophy,” Journal of Molecular and Cellular Cardiology, vol. 29, no. 8, pp. 2213- 2224, 1997/08/01/, 1997.

[0244] [44] S. Rosenkranz, M. Flesch, K. Amann, C. Haeuseler, H. Kilter, U.

Seeland, K.-D. Schluter, and M. Bohm,“Alterations of b-adrenergic signaling and cardiac hypertrophy in transgenic mice overexpressing TGF-bI ,” American Journal of Physiology- Heart and Circulatory Physiology, vol. 283, no. 3, pp. H1253-H1262, 2002.

[0245] [45] A. Brooks, V. Schinde, A. C. Bateman, and P. J. Gallagher,“Interstitial fibrosis in the dilated non-ischaemic myocardium,” Heart, vol. 89, no. 10, pp. 1255-6, Oct, 2003.

[0246] [46] A. Desmouliere, A. Geinoz, F. Gabbiani, and G. Gabbiani,

“Transforming growth factor-beta 1 induces alpha-smooth muscle actin expression in granulation tissue myofibroblasts and in quiescent and growing cultured fibroblasts,” The Journal of Cell Biology, vol. 122, no. 1 , pp. 103-1 1 1 , 1993.

[0247] [47] D. R.P., D. S.T., H. B.P., and T. S.S.W.,“Differential expression of cytokines in the rat heart in response to sustained volume overload,” European Journal of Heart Failure, vol. 6, no. 6, pp. 693-703, 2004.

[0248] [48] F. J. Villarreal, and W. H. Dillmann, “Cardiac hypertrophy-induced changes in mRNA levels for TGF-beta 1 , fibronectin, and collagen,” American Journal of Physiology-Heart and Circulatory Physiology, vol. 262, no. 6, pp. H1861-H1866, 1992.

[0249] [49] A. Calderone, N. Takahashi, N. J. Izzo Jr, C. M. Thaik, and W. S.

Colucci, “Pressure-and volume-induced left ventricular hypertrophies are associated with distinct myocyte phenotypes and differential induction of peptide growth factor mRNAs,” Circulation, vol. 92, no. 9, pp. 2385-2390, 1995.

[0250] [50] M. Pauschinger, D. Knopf, S. Petschauer, A. Doerner, W. Poller, P. L.

Schwimmbeck, U. Kuhl, and H. P. Schultheiss,“Dilated cardiomyopathy is associated with significant changes in collagen type I/Ill ratio,” Circulation, vol. 99, no. 21 , pp. 2750-6, Jun 1 , 1999.

[0251] [51 ] B. L. Riser, P. Cortes, C. Heilig, J. Grondin, S. Ladson-Wofford, D.

Patterson, and R. G. Narins,“Cyclic stretching force selectively up-regulates transforming growth factor-beta isoforms in cultured rat mesangial cells,” The American Journal of Pathology, vol. 148, no. 6, pp. 1915-1923, 1996. [0252] [52] A. J. E. T. van Wamel, C. Ruwhof, L. J. M. van der Valk-Kokshoorn, P.

I. Schrier, and A. van der Laarse,“Stretch-induced paracrine hypertrophic stimuli increase TGF-bI expression in cardiomyocytes,” Molecular and Cellular Biochemistry, vol. 236, no. 1 , pp. 147-153, July 01 , 2002.

[0253] [53] D. J. Milner, G. E. Taffet, X. Wang, T. Pham, T. Tamura, C. Hartley,

M. A. Gerdes, and Y. Capetanaki, “The absence of desmin leads to cardiomyocyte hypertrophy and cardiac dilation with compromised systolic function,” Journal of molecular and cellular cardiology, vol. 31 , no. 1 1 , pp. 2063-2076, 1999.

[0254] [54] Y. Capetanaki, S. Papathanasiou, A. Diokmetzidou, G. Vatsellas, and

M. Tsikitis,“Desmin Related Disease: A Matter of Cell Survival Failure,” Current opinion in cell biology, vol. 0, pp. 1 13-120, 02/1 1 , 2015.

[0255] [55] Elisabeth E. Charrier, A. Asnacios, R. Milloud, R. De Mets, M. Balland,

F. Delort, O. Cardoso, P. Vicart, S. Batonnet-Pichon, and S. Henon,“Desmin Mutation in the C-Terminal Domain Impairs Traction Force Generation in Myoblasts,” Biophysical Journal, vol. 1 10, no. 2, pp. 470-480, 01/19 06/12/received1 1 /23/accepted, 2016.

[0256] [56] A. Heling, R. Zimmermann, S. Kostin, Y. Maeno, S. Hein, B. Devaux,

E. Bauer, W. P. Klovekorn, M. Schlepper, W. Schaper, and J. Schaper, “Increased expression of cytoskeletal, linkage, and extracellular proteins in failing human myocardium,” Circ Res, vol. 86, no. 8, pp. 846-53, Apr 28, 2000.

[0257] [57] J. L. Guichard, M. Rogowski, G. Agnetti, L. Fu, P. Powell, C. C. Wei,

J. Collawn, and L. J. Dell'ltalia, “Desmin loss and mitochondrial damage precede left ventricular systolic failure in volume overload heart failure,” Am J Physiol Heart Circ Physiol, vol. 313, no. 1 , pp. H32-h45, Jul 1 , 2017.

[0258] [58] N. Weisleder, E. Soumaka, S. Abbasi, H. Taegtmeyer, and Y.

Capetanaki,“Cardiomyocyte-specific desmin rescue of desmin null cardiomyopathy excludes vascular involvement,” Journal of Molecular and Cellular Cardiology, vol. 36, no. 1 , pp. 121 - 128, 2004/01/01/, 2004.

[0259] [59] X. Wang, H. Osinska, G. W. Dorn, M. Nieman, J. N. Lorenz, A. M.

Gerdes, S. Witt, T. Kimball, J. Gulick, and J. Robbins,“Mouse Model of Desmin-Related Cardiomyopathy,” Circulation, vol. 103, no. 19, pp. 2402-2407, 2001 .

[0260] [60] E. Takimoto, and D. A. Kass, “Role of oxidative stress in cardiac hypertrophy and remodeling,” Hypertension, vol. 49, no. 2, pp. 241-248, 2007. [0261] [61 ] A. K. Dhalla, M. F. Hill, and P. K. Singal,“Role of oxidative stress in transition of hypertrophy to heart failure,” Journal of the American College of Cardiology, vol. 28, no. 2, pp. 506-514, 1996.

[0262] [62] M. Seddon, Y. H. Looi, and A. M. Shah,“Oxidative stress and redox signalling in cardiac hypertrophy and heart failure,” Heart, vol. 93, no. 8, pp. 903-907, 2007.

[0263] [63] F. J. Giordano,“Oxygen, oxidative stress, hypoxia, and heart failure,”

The Journal of clinical investigation, vol. 1 15, no. 3, pp. 500-508, 2005.

[0264] [64] T. Miinzel, T. Gori, J. F. Keaney Jr, C. Maack, and A. Daiber,

“Pathophysiological role of oxidative stress in systolic and diastolic heart failure and its therapeutic implications,” European heart journal, vol. 36, no. 38, pp. 2555-2564, 2015.

[0265] [65] Z.-W. Liu, H.-T. Zhu, K.-L. Chen, X. Dong, J. Wei, C. Qiu, and J.-H.

Xue,“Protein kinase RNA-like endoplasmic reticulum kinase (PERK) signaling pathway plays a major role in reactive oxygen species (ROS)-mediated endoplasmic reticulum stress- induced apoptosis in diabetic cardiomyopathy,” Cardiovascular diabetoiogy, vol. 12, no. 1 , pp. 158, 2013.

[0266] [66] Z. Ungvari, S. A. Gupte, F. A. Recchia, S. Batkai, and P. Pacher,“Role of oxidative-nitrosative stress and downstream pathways in various forms of cardiomyopathy and heart failure,” Current vascular pharmacology, vol. 3, no. 3, pp. 221 -229, 2005.

[0267] [67] M. Liu, L. Gu, M. S. Sulkin, H. Liu, E.-M. Jeong, I. Greener, A. Xie, I.

R. Efimov, and S. C. Dudley,“Mitochondrial dysfunction causing cardiac sodium channel downregulation in cardiomyopathy,” Journal of molecular and cellular cardiology, vol. 54, pp. 25-34, 2013.

[0268] Example 3

[0269] The disclosed platform is the only known device/chip available that can reproduce the effects of electrical and pressure-volume changes (pressure and stretch) seen in any chamber of the heart at any stage of development. According to aspects, devices, and methods of the present disclosure, electrical stimulation can be precisely coordinated with mechanical loading (onset of systole) to ensure that the electrical impulse is coordinated with cardiomyocyte contractions. Electrical and mechanical stimulation can be applied on 3D cardiac tissue constructs that mimic cellular compositions in embryonic, fetal, neonatal or adult cardiac tissue. Furthermore, pathological conditions including pressure, volume overload and heart failure can be accurately replicated using this device/chip [0270] According to aspects of the present disclosure, can be seamlessly integrated with other organs/tissue to create more complex body-on-a-chip platforms where the cardiac tissue chip acts as a pump to ensure perfusion within the system. Time of day dependent changes in blood pressure, stretch and heart rate can also be reproduced. The cardiac tissue chip can be used to recreate time of day dependent pressure-volume changes (activity and rest) and pressure-volume changes associated with exercise (increased heart rate). Mechanical unloading as seen with mechanical circulatory support (MCS) like left ventricular assist devices (LVADs) can also be reproduced.

[0271] According to aspects of the present disclosure, this chip can be used as a platform to generate mature cardio myocytes (human and other species) by carefully reproducing electrical and mechanical stimulus associated with embryonic heart development to produce structurally and functionally mature cardiomyocytes for research and clinical applications (drug testing, drug discovery and regenerative medicine). This chip can be used to create models of cardiovascular disease including conditions like heart failure, pressure overload and volume overload. This chip can be used to create disease models where interventions (pharmaceutical and device) can be tested. This chip can be used a physiologically relevant model to test drug induced cardiotoxicity of any pharmaceutical drug (off target effects on cardiomyocytes particularly arrhythmogenic potential).

[0272] The cardiac tissue chip can comprise a cell culture chamber and a flow loop, as shown in FIG. 18. The cell culture chamber can comprise an oval shaped well with a thin, flexible membrane bonded to the bottom of the well (see FIGs. 2A-2B; FIG. 20). Two posts are mounted on the thin membrane for anchorage and suspension of the 3D engineered cardiac tissue.

[0273] Within the cell culture chamber, a 3D construct comprising cardiac cells suspended within a hydrogel is cast using a removable mold around the two posts (an embodiment of a mold is shown in FIG. 19 in position with the two posts). The hydrogel can be made of any material that supports cardiac cell survival including but not limited to fibrin, collagen, matrigel, alginate or a combination of different hydrogels. Once cells adapt to the gel, the gel compacts and forms a 3D tissue that is suspended between the posts.

[0274] The engineered tissue within the chamber can be integrated within a flow loop. The flow loop comprises a medium reservoir, one-way flow control valves, resistance element and a programmable pressure generator. The mechanism of action is based on establishing dynamic pressure differences above and below the cell culture chamber membrane in order to reproduce the cardiac cycle. Placing the media reservoir at a specified height creates a downward filling pressure (end-diastolic pressure) on the cell culture member to mimic a diastolic filling pressure. The programmable pneumatic pump is used to increase the pressure below the cell culture membrane in order to return the membrane to its original position and generate a systolic (ejection) pressure. The systolic pressure can be set by adjusting the resistance element. The cycling of the pneumatic pump coupled with one-way flow control values on either side of the cell culture chamber allows the actuation of the culture membrane to displace fluid in one direction with each contraction.

[0275] Stretch can be accomplished by placing a curved support below the flexible membrane to direct stretch as shown in FIGs. 5C and 5D. As the chamber fills the center of the membrane is constrained by the support causing the edges to move downward and pull the posts outward. These results in imposition of stretch on the 3D tissue suspended between the posts.

[0276] Various conditions including embryonic, adult, diseased, interventions, exercise etc can be imposed on cells within the 3D construct by adjusting the height of the reservoir, the level of stretch of the thin membrane, applied pressure from the pressure generator, use of additional pumps to unload the chamber (MCS/LVAD).

[0277] Advantages and improvements over prior devices include that this is the only platform where the pressure-volume changes seen in any chamber of the heart at any stage of development can be accurately recreated. A combination of electrical and mechanical loading in a single platform imposes these stresses in a physiologically relevant manner. It is the only platform known to date that can accurately recreate hemodynamics associated with development and disease. It is the only known platform to date where interventions like MCS/LVAD can be evaluated. Applications of aspects of the present disclosure include: generation of mature cardio myocytes; disease models; exercise models; time of day dependent changes; drug feasibility testing; drug toxicity testing; and evaluation of pharmaceutical and device-based interventions.

[0278] Example 4

[0279] Current State-of-the-art: Maturation of hiPS-CMs is a critical problem that needs to be addressed. While there have been several efforts to mature hiPS-CMs, the approach that has shown the most promise is the use of electromechanical‘intensity training’ applied to early stage hiPS-CMs as described by Ronaldson-Bouchard et al. (Ronaldson- Bouchard, K. et al. Advanced maturation of human cardiac tissue grown from pluripotent stem cells. Nature 556, 239-243, doi:10.1038/s41586-018-0016-3 (2018), herein incorporated by reference in its entirety, especially regarding iPS-CM differentiation protocol, although the skilled artisan would understand that other differentiation protocols may be suitable) which achieved superior structural and functional maturation in terms of sarcomeric structure, ultrastructural morphology (t-tubules), gene expression, calcium handling, inotropic response, metabolism and mitochondrial numbers. This approach however fell short in terms of electrophysiological maturation and force generation. The most important findings of this paper are: (1) hiPS-CMs respond to electromechanical stimulation via maturation especially if the intensity of the stimulation is gradually increased and (2) Early during the differentiation process (days 12-20) there is a window of cellular plasticity that can be exploited to achieve maturation following which hiPS-CMs become terminally differentiated.

[0280] Shortcomings of the‘intensity training’ approach: The main shortcomings of this approach are the following: (1 ) The setup used provides passive resistance (flexible posts) to electrically stimulated tissue contractions. This approach does not provide preload (i.e.) mimic stretch and filling pressure as seen during active filling of the ventricular chamber (diastole) or systolic pressure seen during isovolumetric contractions. Multiple avian embryonic heart development studies suggest that pressure is a critical stimulus that ensures proper heart development and an increase or decrease in developed pressures results in congenital heart defects. (2) This protocol mimics only a certain window of embryonic heart development (FIG. 22, red) (i.e.) the period immediately following node (pacemaker) formation where the heart rate increases rapidly to maintain cardiac output (CO). In the growing embryo the CO needs to increase as the embryo grows in size to ensure survival. When the heart is small, CO (stroke volume (SV) x heart rate (HR)) is increased via a rapid increase in HR. However, after the heart chambers increase in size, the SV increases and HR gradually drops. Subsequent increase in CO is primarily accomplished via an increase in ventricular chamber size (SV). This protocol does not mimic the period prior to node formation (FIG. 22, green) where cardio myocytes beat spontaneously to induce peristaltic flow and experience small changes in pressure and stretch or the period where the HR begins to decrease gradually and the stretch or preload (due to increased SV) increases (FIG. 22, blue).

[0281] Novelty of the Proposed Approach with the BCTM: The BCTM accurately mimics stretch and end-diastolic pressure seen during the filling phase and systolic pressure during isovolumetric contractions (FIG. 5E). The BCTM is therefore an accurate mimic of hemodynamic loading and unloading of the heart. Using the BCTM pressure and stretch due to peristaltic flow in the linear heart tube can be more accurately mimic’d, rapidly increasing electromechanical stimulation following establishment of nodes and finally gradually increase pressure and stretch to post-natal and adult levels similar to what occurs during human heart development (FIG. 22).

[0282] Example 4

[0283] Mechanical Circulatory Support (MCS), especially left ventricular assist devices (LVADs) are now standard therapy for patients with HF (HF) that is refractive to medical therapy. LVADs placed in parallel to the left ventricle (LV) cause pressure and volume unloading of the LV and restore CO and organ perfusion. While almost all patients with LVADs experience an improvement in cardiovascular function, a subset of patients also experience an improvement in myocardial function suggesting that LVADs can potentially be used as a bridge to recovery. The improvement in myocardial function is possibly due to regression of cardiac hypertrophy (via cardiac atrophy) in response to LVAD ventricular unloading. Despite the promise of myocardial recovery with MCS, the mechanisms via which hypertrophy is reversed and myocardial function is recovered remains poorly understood and may require comprehensive evaluation of biological processes involved at the cellular level. Physiologically relevant Cardiac Tissue Chips to study the effects of mechanical unloading of the heart on myocardial recovery require the ability to (1 ) Generate pathologically hypertrophic tissue under conditions of pressure overload (PO) that ultimately leads to HF and (2) The ability to subject pathologically hypertrophic tissue to mechanical unloading similar to unloading seen with LVADs. Such a model can be used to determine if unloading of pathologically hypertrophic tissue can promote beneficial remodeling. The BCTM can reproduce both pathological hypertrophy and unloading seen with LVADs. According to the present example, the BCTM can be used to generate pathologically hypertrophic cardiac tissue associated with PO and HF to determine if mechanical unloading can promote beneficial reverse remodeling.

[0284] Preliminary Studies: Data was recently generated that clearly demonstrates that cardiac tissue constructs composed of cardiomyocytes and cardiac fibroblasts when subjected to conditions of PO (160 mmHg systolic pressure) results in concentric hypertrophy and fibrotic remodeling of the engineered tissue (FIGs. 13A-13B). When the engineered tissue was subjected to PO in the presence of TGF-b, the amount of collagen I deposition was significantly increased in comparison to engineered tissue subject to normal pressure or tissue cultured with TGF-b under static conditions (FIGs. 13C-13E). Other hallmarks of pathological hypertrophy including concentric hypertrophy and increased pro-fibrotic and pro- inflammatory gene expression ( Coll , Col3, TGF-b, Timpl, Timp2, SOD1, SOD2) were also observed. The BCTM setup was also modified to mimic left ventricular unloading with LVAD support. The setup comprises a peristaltic pump that unloads the cell culture chamber to mimic unloading seen with commonly used continuous flow LVADs (FIG. 24A). Unloading the BCTM chamber set to mimic HF conditions results in a decrease in preload as the chamber filling is incomplete due to the continuous flow pump, and reduced end-diastolic pressure (EDP) and systolic pressure (SP) (FIG. 24B), closely mimicking changes seen in patients.

[0285] Research Design and Methods: Engineered cardiac tissue can be generated representative of patients with HF that are candidates for LVAD placement. Engineered tissue can be composed of hiPS-CMs and cardiac fibroblasts (3: 1 ratio) seeded in fibrin gel. Constructs will be maintained in static culture for 2 days until the fibrin gel compacts and then integrated into the BCTM flow loop and gradually subject to pressure and stretch over a period of 4 days until the constructs experience normal adult pressure-volume (PV) loading (120 mmHg SP, 0-10% stretch, 10 mmHg EDP) and electrical pacing. Prior work shows that it is important to gradually impose stress on hiPS-CMs to ensure cell survival. Typically, cardiac tissue remodeling occurs under PO due to conditions like hypertension or aortic stenosis and then gradually progresses to HF. PO represents a situation with abnormally high systolic SP (> 150 mmHg) but end diastolic volume (EDV) and EDP remains at physiological levels. As PO progresses towards HF, contractile function is compromised resulting in a gradual drop in systolic pressure and an increase in both residual and EDV and an increased EDP (> 15-20 mmHg EDP). Using the BCTM, engineered tissue cultured under normal PV loading can be transitioned to conditions of PO (180 mmHg SP, 0-10% stretch, 10 mmHg EDP) in the presence of 2 ng/mL of TGF-b, which is sufficient to observe extensive fibrosis and collagen I deposition. Following 5 days in culture under PO, the loading conditions can be gradually changed to mimic HF (120 mmHg SP, 10-20% stretch, 25 mmHg EDP) over a period of 4 days following during which engineered tissue under conditions of HF will be maintained or provided LVAD support using the pump for an additional 5 days. To more closely mimic observations previous observations with infiltration and M1 polarization of macrophages in human and animal HF, human monocytes (human monocytes from

StemCell Technology®) can be introduced into the perfusing medium during conditions of PO to evaluate infiltration and polarization of monocyte/macrophage populations during the transition from PO to HF. Constructs can be evaluated prior to LVAD support to confirm pathological hypertrophic remodeling and after LVAD support to determine if unloading promotes myocardial recovery. Polarization of infiltrated monocytes/macrophages will also be evaluated before and after LVAD support to evaluate the role of unloading on immune cell polarization and myocardial recovery. Engineered constructs under normal PV loading can be used as a control.

[0286] Example 5

[0287] FIGs. 25A-25C: FIG. 25A is a re duced-to- practice embodiment of a complete BCTM flow loop to subject engineered cardiac tissue to pressure-volume loading; FIG. 25B is a schematic and actual image of the 3D muscle strip integrated within the BCTM flow loop chamber where pressure-volume changes can be imposed. FIG. 25C are graphs showing examples of putative changes from normal pressure-volume loading to pathological conditions like pressure and volume overload.

[0288] An aspect of technological innovation in this example and disclosure is the Biomimetic Cardiac Tissue Model (BCTM) and its ability to culture in vivo -like 3D cardiac tissue under electrical and mechanical stimulation similar to how electrical impulses result in pressure-volume changes associated with hemodynamic loading and unloading (FIGs. 25A- 25C). Devices, systems, and methods of the present disclosure have the following advantages that meet National Center for Advancement of Translational Sciences (NCATS) criteria for Tissue Chips: (1 ) Multi-cellular 3D architecture associated with in-vivo cardiac tissue, (2) Functional representation of complex hemodynamic loading and unloading (pressure-volume changes) associated with normal and pathological tissue, (3) Reproducible and long-term maintenance of engineered constructs within the BCTM for up to 30 days in culture without compromise in cell viability or function. (4) Accurate representation of developmental, normal and disease states via control of hemodynamic loading and pro- inflammatory stimuli.

[0289] Unless defined otherwise, all technical and scientific terms used have the same meaning as commonly understood by one of ordinary skill in the art to which this disclosure belongs. Although any methods and materials similar or equivalent to those described can also be used in the practice or testing of the present disclosure, the preferred methods and materials are now described.

[0290] Embodiments of the present disclosure will employ, unless otherwise indicated, techniques of separating, testing, and constructing materials, which are within the skill of the art. Such techniques are explained fully in the literature.

[0291] It should be emphasized that the above-described embodiments are merely examples of possible implementations. Many variations and modifications may be made to the above-described embodiments without departing from the principles of the present disclosure. All such modifications and variations are intended to be included herein within the scope of this disclosure and protected by the following claims. As used herein,“about” means ± 5% or ±10% of the reference value.