Tissue engineering three-dimensional cardiac organoids provides the potential to develop an alternative treatment solution for congenital cardiac conditions such as hypoplastic left heart syndrome (HLHS).1 , 2 Hypoplastic left heart syndrome presents with multiple structural cardiac defects, which include, but are not limited to left ventricle and associated valve underdevelopment and stenoses; a replacement engineered construct must possess the properties of tissue-type architecture, with inducible coherent contractions, propensity to deform in response to systolic and diastolic loads, and action potential propagation syncytia.2 Current HLHS treatments are designed to surgically refunctionalize the right ventricle to adapt to the role of the left ventricle, through a three-stage process: developing an aorta linked to the right ventricle and reconfiguring returning blood flow for direct passage to the lungs for reoxygenation.1 Directly replacing the underdeveloped left ventricle component of HLHS presents an optimal treatment solution compared with current right ventricle refunctionalization strategies; thus, bioengineering a complete three-dimensional left ventricle with trileaflet valves may provide a foundation to develop a direct tissue engineered solution. The ability to bioengineer such a model was tested in the current study by valve development, total ventricle assembly, and functional assessment.
Underdevelopment of both the left ventricle and aortic valve are critical presentations in HLHS; current treatment strategies focus on refunctionalizing the right ventricle. The ability to supersede left ventricle chamber and valve pathophysiologies by means of tissue engineering has been studied independently, for application as either cell delivery platforms or total replacement structures.3–9 Cardiac tissue engineered advances have produced cell-based chambers and micropumps.10–12 Costa et al. 8 and Ott et al. 9 developed simplified organoid chambers and structurally preformed cardiac constructs, respectively, exhibiting primary ventricular pump function. These studies focused on completing the three-dimensional bioengineered open ventricle (BEOV) by development of a tissue engineered trileaflet valve, as aortic valve stenosis and left ventricle chamber underdevelopment are the primary pathophysiologic presentations in HLHS.
We have developed and characterized a BEOV and optimized the chamber with respect to cell delivery and retention and function using bioreactor conditioning. Molds were designed to emulate the geometry of a human neonate left ventricle chamber; scaffolds were fabricated from chitosan based on previous characterization studies.6 , 13 , 14 First, we investigated an effective two-stage cell delivery system to improve cell retention.15 Subsequently, we studied the effect of pulsatile flow stretch conditioning on BEOV function with respect to electrophysiologic syncytia and pressure generation.16
All protocols were approved by the Institutional Animal Care and Use Committee (IACUC) in accordance with the “Guide for Care and Use of Laboratory Animals” (NIH Publication 86-23, 1986)
Bioengineered Complete Ventricle Scaffold Fabrication and Preparation
Bioengineered open ventricle molds were fabricated according to a previously described method.13 Computer-aided design (CAD) software and three-dimensional printing were used to develop an open ventricle, with the same geometry as a neonatal left ventricle. Bioengineered trileaflet valve (BETV) molds were fabricated using CAD software and computer numerical code (CNC) machining, with the same geometry as a human infant aortic trileaflet valve (Figure 1A, i). Scaffolds for both the BEOV and BETV were fabricated according to a previously described method13; chitosan (Carbomer Inc., San Diego, CA) was solubilized in 0.2 M acetic acid (Glacial Acetic Acid; Macron Fine Chemicals, Centre Valley, PA) in deionized water and transferred into the molds. Filled molds were frozen at −80°C for 24 hours and then lyophilized for 24 hours (Labconco Freezone 4.5; Labconco Corporation, Kansas City, MO) (Figure 1B, i).
Lyophilized BETV were secured into the lyophilized BEOV using cyanoacrylate adhesive and polypropylene sutures (AD Surgical, Sunnyvale, CA) (Figure 2A, i and ii); BETV were attached in a closed state, to produce the bioengineered complete ventricle (BECV) model.
Lyophilized scaffolds were rehydrated in preparation for cellularization by sequential immersion in 0.1 M sodium hydroxide (NaOH) (NaOH pellets, Macron Fine Chemicals) in deionized water for 30 minutes, Phosphate Buffered Saline (PBS) washing, sterilization in 80% ethanol for 1 hour, and rinsing with sterile PBS. Rehydrated scaffolds were then exposed to ultraviolet (UV) light for 20 minutes and air dried in a laminar flow hood.
Isolation of Primary Cardiac Myocytes and Fabrication of Artificial Heart Muscle
Cardiac cells were isolated from the hearts of 2–3-day-old neonatal Sprague-Dawley rats using an established method.17
Artificial heart muscle (AHM) was fabricated using an established protocol.18 Thirty-five millimeters tissue culture plates were coated with 2 ml of SYLGARD (PDMS, type 184 silicone elastomer) (Dow Chemical Corp., Midland, MI) and air dried for 2 weeks. Plates were sterilized in 80% ethanol, after which 4 minutien pins of 0.1 mm diameter (Fine Science Tools, Foster City, CA) were anchored into the PDMS gel, in a 2 cm × 2 cm square. The fibrin gel was formed by plating 1 ml of 10 U/ml thrombin in culture media, composed of M199 (Life Technologies, Grand Island, NY) with 20% F12k (Life Technologies), 10% foetal bovine serum, 5% bovine serum, 1% antibiotic-antimycotic, 40 mg/ml hydrocortisone, 100 ng/ml insulin, and 2 mg/ml amino-caproic acid, to inhibit fibrinolysis, followed by 500 µl of 20 mg.ml of fibrinogen in saline; plates were shaken to ensure complete mixing and incubated for 30 minutes to promote gel formation. Isolated primary cardiac cells were suspended at 2 million cells/ml in culture media, and 2 ml of the cell solution was added to each plate after gel formation. At 3–4 days of culture in 37°C supplied with 5% carbondioxide (CO2), cells organized, resulting in spontaneous contraction for gel delamination to produce AHM, which formed around the minutien pin anchor points.
Bioengineered Complete Ventricle Cellularization and Culture
The prepared BECV scaffolds were secured on cotter pins and mounted onto individual platforms (Figure 2B, i). Constructs were suspended from the top of the platform, to increase the surface area in contact with the culture medium and facilitate cell viability during the culture period.
The BECV chamber was cellularized as per a previously described method.13 Stage one of BECV chamber cellularization is direct cell injection (DCI) of isolated primary cardiac cells (Figue 2B, ii). DCI was performed using a 16 gauge hypodermic needle (BD, Franklin Lakes, NJ). A total of 1 × 107 rat neonatal cells, suspended in 1 ml of culture media, were injected into the scaffold, through the outer scaffold surface, at multiple points.
Stage two of BECV chamber cellularization is immediate wrapping of an AHM patch around the stage one cellularized construct (Figure 2B, ii) and securing with polypropylene sutures (AD Surgical, Sunnyvale, CA). The AHM was wrapped proximal to the scaffold outer surface, with cardiac cells between the scaffold and fibrin gel.
Cardiac fibroblasts were passively seeding onto the BECV, by deposition into the valve construct. A total of 6 × 105 rat neonatal fibroblast cells, suspended in 1 ml of culture media, were passively seeded on the BECV scaffold.
Platform-mounted cellularized constructs were secured into individual cubicles, in a custom bioreactor housing for perfusion culture (Figure 2C, i). A total of 350 ml of culture media was added to the bioreactor, to submerge the constructs; a continuous unidirectional flow of media was maintained at a rate of 50 ml/min using a Masterflex L/S pump (Model 77200–62; Cole-Palmer, Vernon Hills, IL). Each cubicle contained 100 ml of culture media, with a flow rate of 16 ml/min; the total volume of media in the tubing was 50 ml. Constructs were maintained in an incubator at 37°C supplied with 5% CO2 for 3 days. The perfusion flow rate was adjusted to be equivalent to that of the high flow rate described by Radisic et al. 19 as a function of relative scaffold volume and cell load.
The outcome model constructs are referred to as the BECV for the remainder of the study.
Bioengineered Complete Ventricle Pressure Measurement
Bioengineered complete ventricle pressure was measured according to a previously described method.16 Bioengineered complete ventricle pressure was measured using a Mikro-Tip Catheter transducer (SPR-524; Millar Instruments, Houston, TX); the catheter was placed through the upper flow port in the bioreactor platform and through the closed valve system into the BECV chamber. The catheter was connected to a quad bridge amplifier (FE224; AD Instruments, Dunedin, NZ), and data were acquired through a 16 channel PowerLab system (PL3516/P; AD Instruments). LabChart was used for data analysis, with the peak analysis module, to calculate average pressure generated.
Bioengineered Complete Ventricle Biopotential Measurement
Cardiac biopotential measurements were obtained to assess BECV cardiac myocyte electrophysiology and activity. Biopotential activity was measured using a 32 electrode biopotential sensor, according to previously described methods.13 , 15 , 16
The biopotential activity of each cell delivery model was measured for a total of 60 seconds. The acquired data were analyzed using RND-m filed (open source) provided by Intan Technologies, to import the data into MATLAB. Raw data channels were examined for periodic waveforms in multiple channels, within a specific time interval and processed using “crosscorr” (MATLAB signal processing toolbox), to determine the average time delay and conduction velocity for each channel, or electrode, with respect to the remaining 31 channels.
After the 3-day perfusion culture period, BECV were prepared to produce planar tissue sections of 20 µm thickness. Bioengineered trileaflet valve and BECV were separated; constructs were placed in a peel-a-way disposable embedding molds (VWR International, Radnor, PA) immediately after the 3-day culture period. Constructs were immersed in Tissue-Tek optimum cutting temperature compound (VWR International) and frozen at −80°C for 24 hours. Planar sections of 20 µm thickness were cut using a Cryotome (ThermoScientific, Waltham, MA) and placed onto VWR microslides for staining. Immunohistochemistry was used to determine cell distribution and characteristics in BECV, for the factors, cardiac troponin I, a-actinin, and connexin43 (Cx43) in the BECV chamber, and Vimentin for fibroblasts.
Nonspecific epitope antigens were blocked with 10% goat serum in 0.05% Triton-X100 PBS, at room temperature for 1 hour. To show contractile factors, sections were incubated with rabbit antitroponin I, 1:100 (ab47003; Abcam, Cambridge, MA). To show sarcomere presence, sections were incubated with mouse anti-a-actinin, 1:200 (Sigma-Aldrich, St Louis, MO, Catalog No. A7811).
To show cell-cell interactions, sections were incubated with rabbit anti-Cx43, 1:100 (ab11370, Abcam). To show fibrobast presence, sections were incubated with mouse anti-Vimentin, 1:5000 (ab195878). All sections were counterstained for nuclei with 4,6-diamidine-2-phenzylindole (DAPI) (2.5 µg/ml) for 5 minutes at room temperature. Fluorescent images were obtained with a Nikon C2+ confocal laser scanning microscope (Nikon Instruments Inc., Melville, NY).
Fabrication of the Bioengineered Trileaflet Valve
The ability to develop a BECV was first addressed by the design and fabrication of a BETV, to compete the previously developed BEOV model. The human aortic trileaflet valve has a variable inlet diameter (22–28 mm) and is comprised of three geometrically distinct leaflets (Figure 1A, i). A physiologically simulative BETV mold was designed using CAD software and CNC machining, (Figure 1A, ii and iii); the BETV model accounts for geometrical distinction between individual leaflets.
The BETV scaffold was produced at a concentration in keeping with the BEOV chamber, of 2.5% w/v chitosan in acetic acid (Figure 1B, i). The lyophilized BETV scaffold (Figure 1B, ii) retained a three-dimensional self-supporting geometry structure, congruent with the geometry, and shape of the designed mold. Lyophilization is a process wherein polymer fibers orient around ice crystals, which are subsequently removed by sublimation through application of a vacuum, to leave a pore; the observable striations are attributed to this process. Rehydration in preparation for cellularization results in the loss of observable surface striations (Figure 1B, iii) as a result of material swelling and a subsequent reduction in pore size.
The CAD BETV model was approximately 30% smaller compared with the adult human aortic valve geometry (Figure 1C, i)20 to mimic a neonate structure; all proportions are described relative to the inflow diameter, d. The uniformity and reproducibility of the CNC mold and scaffolds were assessed relative to the CAD model (Figure 1C, ii). The CNC mold (n = 3) showed significant variability compared with the CAD model with respect to height (CAD versus CNC; CNC 8% smaller versus CAD) and L2 radius and thickness (CAD versus CNC; CNC 60% greater versus CAD); these outcomes may be attributable to the complexity of the CAD model coupled with machining limitations. Bioengineered trileaflet valve scaffolds each displayed significant variability with an increase in leaflet radii for both the lyophilized (n = 10) and rehydrated (n = 10) samples, compared with the CAD model ranging from 6–50% (CAD versus lyophilized BETV) and 1430% (CAD versus rehydrated BETV). In lyophilized scaffolds, the length of the open leaflet was 11% smaller compared with the CAD model; the inflow diameter was 28% narrower in the rehydrated scaffolds.
Sodium hydroxide rehydration of the lyophilized scaffolds resulted in a significant narrowing and lengthening of the BETV (lyophilized versus rehydrated BETV, 25% reduced inflow diameter, 6% increased length); relative leaflet radii remained similar; however, a reduction in leaflet thickness was noted. The observed and measured effects of rehydration may be attributed to NaOH acetylation with chitosan polymers, resulting in cross linking.
Fabrication of the Bioengineered Complete Ventricle
The development of the BECV by assembly of the BEOV and BETV was assessed from CAD design to chitosan scaffold (Figure 2A). Individual components were secured in place to seal and complete the chamber using cyanoacrylate adhesive and sutures, respectively. Bioengineered complete ventricle scaffold processing varied significantly with respect to distance from valve peak to BEOV inner face apex; both lyophilized and rehydrated BECV scaffold (n = 3 for each) distances were significantly lower than that of the CAD model (rehydrated BECV scaffold distance = 10.61 ± 0.18 mm; CAD versus lyophilized/rehydrated BECV scaffolds, p < 0.05 for all, t-test). The inner cavity height was differed significantly between the lyophilized and rehydrated BECV scaffolds; the increased length measured in the rehydrated scaffold can be attributed to NaOH acetylation, potentiating the formation of cross-links as characterized on our previous studies.13
The rehydrated BECV scaffold retained its shape and frame (Figure 2B, iv). On attachment to a custom-fabricated suspension platform, the chamber was cellularized as per the two-stage method described in our previous studies (Figure 2B, i and ii).15 Valve cellularization was a passive process; cardiac fibroblasts were isolated from a mixed population of Sprague-Dawley rat neonatal cardiac cells and suspended in culture medium at a concentration of 6 × 105 cells/ml. One milliliter of cell solution was delivered onto the BETV to immerse all valve leaflets. The cellularized BECV retained their shape and frame on cellularization and in the bioreactor during the culture period (Figure 2B, iv and C).
Bioengineered Complete Ventricle Pump Function
In total, three BECV were analyzed in this study; all samples displayed macroscopic “twitch-type” function at cellularization (see Video, Supplemental Data 1A, https://youtu.be/K7ab3aiuuF0). Observed contractions became smoother and more pervasive after conditioning (see Video, Supplemental Data 1B, https://youtu.be/K7ab3aiuuF0). Bioengineered complete ventricle pressure output (Figure 3) validates the macroscopic observations; pressure generation after conditioning either remained comparable to that of the BECV at cellularization or was significantly increased (Figure 3; BECV1 at cellularization versus BECV1, p < 0.05, t-test) indicating that the force generated at each contraction remained relatively consistent after conditioning. The pressure generated both at cellularization and after conditioning remained significantly lower than the left ventricle end-diastolic pressure in both rats (18 ± 1 mm Hg)21 and human neonates (11 ± 3 mm Hg)22 (p < 0.05 for all, t-test). The increased pressure reading in BECV1 after conditioning may be attributed to the formation of a tight seal at BECV assembly to minimize leaks and the variability associated with the two-stage cellularization method.
Bioengineered Complete Ventricle Cardiac Biopotential Activity
Biopotentials represent cardiac myocyte depolarization and repolarization activity; this activity was measureable for all three BECV constructs, and the output was compared with our previous two-stage cellularized open chamber (two-stage perfusion cellularized ventricle [TPCV]) model, which was cultured under the same conditions as the BECV.15 Bioengineered complete ventricle raw biopotential data were examined for periodic waveforms using MATLAB to produce bioplots (Figure 4A, upper). Optical maps were formulated to observe impulse propagation in the BECV (Figure 4A, lower); both maps show individual reference channels compared with the remaining 31 channels. Overall, impulse propagation time delays appear lower in BECV (Figure 4A, lower left) compared with the TPCV (data not shown),15 with time delay uniformity in the reference channels; in particular, a consistently lengthier delay for channels 4, 6, 8, 17, 24, 31, and 32.
Conduction velocity was calculated using the known distance travelled between channels (Figure 4A, lower right). On average, the total conduction velocity was 43.9 ± 49.9 cm/s (mean ± standard deviation) for the BECV and 22.1 ± 30.2 cm/s for the TPCV.15 Bioengineered complete ventricle conduction velocity was significantly accelerated compared with the TPCV (TPCV versus BECV, p < 0.05, t-test); however, there was significant variability between individual channels in both models (Figure 4B).15
The distribution and average biopotential amplitudes were calculated over the total measurement period for the BECV (Figure 4C). Bioengineered complete ventricle displayed a more consistent biopotential amplitude output (1022.9 ± 418.4 µV, mean ± standard deviation), compared with the notably variable open chamber TPCV model (1731.1 ± 1827.5 µV).15 The notably lower BECV biopotential amplitude may be associated with the significantly increased conduction velocity observed.
Characterization of Bioengineered Complete Ventricle
Structurally, the BECV chamber is composed of syncytial-type aggregates of cardiac myocytes, present at the chitosan-AHM patch boundary (Figure 5A). Furthermore, cardiac myocytes were diffusely present within the BECV chamber scaffold and therefore evidenced few striated sarcomeres. Both the contractile protein, cardiac troponin I (CTnI), and the gap junction protein, connexin 43 (Cx43), were well established at syncytial points occurring at the scaffold-AHM boundary, to validate the contractile (see Video, Supplemental Data 1, https://youtu.be/K7ab3aiuuF0; Figure 5A) and biopotential outputs (Figure 4A) observed. The BECV chamber demonstrates the potential to produce a structure resembling cardiac myofibril-type organization and sarcomeric structure.8 , 23 The BECV valves exhibited the presence of randomly oriented, diffusely distributed cardiac fibroblasts (Figure 5B) to indicate the potential of the engineered valve to ameliorate fibroblast infiltration and retention through passive cell seeding.
This study describes the development of a novel BECV by fabrication of a BETV, for assembly with a BEOV. The BECV replicates the geometry of the human neonate left ventricle, with valve, to create a closed system; the chitosan scaffold is a three-dimensional self-support structure. The BECV chamber was cellularized using the two-stage approach of DCI, followed by AHM wrapping. The BECV was produced after unidirectional continuous perfusion culture for 3 days; the model generated spontaneous pump function with appropriate biopotential activity.
Cell-based chambers and micropumps represent current advances in cardiac tissue engineering applicable to the development of an engineered left ventricle.10–12 Simplified organoid chambers, wherein the system is closed by attachment to a Langendorff perfusion system, 8 and preformed complete cardiac chambers9 studies describe the use of differential pressure loads to maintain chamber structure and recreate systolic and diastolic pressure loads to ameliorate ventricular-type properties. Present three-dimensional ventricle studies demonstrate a niche in the development of a complete bioengineered left ventricle with trileaflet valve, of self-supporting structure and spontaneous ventricle-type function.
Our previous work describes the process to engineer a BEOV,13 optimize cellularization in an open BEOV chamber,15 and mechanically stretch condition an open chamber BEOV for inherent pump function.16 Primarily, the BEOV study focused on open chamber design and fabrication, with identification and characterization of a suitable, biocompatible scaffold biomaterial; direct cardiac cell injection resulted in cell-scaffold interactions detected by Fourier transform infrared spectroscopy and biopotential activity; however, cell retention remained low. An effective two-stage open chamber cellularization model (TPCV) was formulated to improve cell retention, with direct cardiac cell injection followed by AHM wrapping of the scaffold; cell retention and biopotential activity improved in this model; however, pressure generation was identified to be an integral component in ventricle design.
The BECV demonstrated improved functional and histologic characteristics in comparison with our previously modeled open TPCV system.15 Bioengineered complete ventricle elicited measurable output pressure (Figure 3), with increasingly pervasive contractile function over time (see Video, Supplemental Data 1, https://youtu.be/K7ab3aiuuF0). The pressure generated both at cellularization and after conditioning remained significantly lower than the left ventricle end-diastolic pressure in both rats (18 ± 1 mm Hg)21 and human neonates (11 ± 3 mm Hg)22 (p < 0.05 for all, t-test). Conduction velocity was approximately 50% faster than that of the open TPCV system (Figure 4) (BECV versus TPCV, p < 0.05, t-test); the final conduction velocity remains comparable with that of other in vitro cardiac engineered tissue constructs.24 Bioengineered complete ventricle displayed a more consistent biopotential amplitude compared with the TPCV (BEOV versus TPCV, p < 0.05, t-test); the output values were within the reference ranges for rats (0.495–0.775 mV)25 and humans (0.1–0.5 mV).26 At syncytial points, cell aggregates were observable at the AHM boundary with some sarcomeric presence and cardiac myocyte organization presence; cardiac myocyte infiltration throughout the BECV chamber was diffuse. The presence of these histologic characteristics validates contractile (see Video, Supplemental Data 1, https://youtu.be/K7ab3aiuuF0), pressure (Figures 3 and 5A), and biopotential outputs (Figure 4).
This novel, complete bioengineered ventricle provides a platform on which numerous optimizations can be implemented to improve functional properties. Material characteristics demonstrate the ability to retain the designed geometries and architecture, with biocompatibility to enable cell infiltration and retention. The mechanical properties of the BECV, with regards to pressure generation in the current study, are still subsidiary to those of the native human neonate left ventricle (11 ± 3 mm Hg).22 Bioengineered complete ventricle electrophysiologic activity is comparable with human neonate outputs (0.1–0.5 mV)25; however, biopotential syncytial properties remain significantly variable. The development of a novel bioengineered left ventricle, complete with trileaflet valve, which generates spontaneous pressure and biopotential output with simple perfusion culture, suggests that further optimization of cell delivery and multistimulus bioreactor conditioning may lead to an engineered ventricle comparable with that of the native tissue.
Our BECV model represents an early attempt to bioengineer a complete bioartificial ventricle that has the future potential to benefit patients with HLHS. Our current configuration of the ventricle highlights many of the key technological components, inclusive of anatomically matched geometry, regulation of fluid flow through bioengineered valves, and cellularization using primary neonatal cardiac myocytes. However, many scientific and technological challenges remain prior to the utilization of BECVs in HLHS patients. Before clinical utilization, the current BECV would require the addition of inflow-outflow valves; efforts are currently ongoing in our laboratory. Furthermore, the current BECV model relies upon primary cardiac myocytes isolated from neonatal rat hearts for cellularization. Clinical use of the BECV will necessitate a different cell source; for example, placental mesenchymal stem cells that have been reprogrammed to form cardiac myocytes. Although many challenges remain, the BECV model represents an early attempt toward the fabrication of ventricles; upon further development, BECVs have the long-term future potential to be applied in a clinical setting to help HLHS patients.
The fabrication of this novel BECV, coupled with the results of the current proof-of-concept study, may aid in further developing a functional engineered left ventricle for clinical application in HLHS. Additionally, the presented model may be used to study the effects of differential stimuli on tissue development, in disease and injury models and as a high throughput screening system, as it provides a three-dimensional simulative physiologic platform, comparable with that of the in vivo environment.
1. Barron DJ, Kilby MD, Davies B, Wright JG, Jones TJ, Brawn WJ. Hypoplastic left heart syndrome
. Lancet 2009.374: 551–564.
2. Zimmerman WH, Schnierderbanger K, Schubert P, et al. Tissue engineering of a differentiated cardiac
muscle construct. Circ Res. 2002.90:223–30.
3. Neuenschwander S, Hoerstrup SP. Heart valve tissue engineering. Transpl Immunol 2004.12: 359–365.
4. Sutherland FW, Perry TE, Yu Y, et al. From stem cells to viable autologous semilunar heart valve. Circulation 2005.111: 2783–2791.
5. Dohmen PM. Tissue engineered aortic valve. HSR Proc Intensive Care Cardiovasc Anesth 2012.4: 89–93.
6. Albanna MZ, Bou-Akl TH, Walters HL 3rd, Matthew HW. Improving the mechanical properties of chitosan-based heart valve scaffolds using chitosan fibers. J Mech Behav Biomed Mater 2012.5: 171–180.
7. Mendelson K, Schoen FJ. Heart valve tissue engineering: Concepts, approaches, progress, and challenges. Ann Biomed Eng 2006.34: 1799–1819.
8. Lee EJ, Kim DE, Azeloglu EU, Costa KD. Engineered cardiac
organoid chambers: Toward a functional biological model ventricle. Tissue Eng Part A 2008.14: 215–225.
9. Ott HC, Matthiesen TS, Goh SK, et al. Perfusion-decellularized matrix: Using nature’s platform to engineer a bioartificial heart. Nat Med 2008.14: 213–221.
10. Gonen-Wadmany M, Gepstein L, Seliktar D. Controlling the cellular organization of tissue-engineered cardiac
constructs. Ann N Y Acad Sci 2004.1015: 299–311.
11. Khait L, Birla RK. Cell-based cardiac
pumps and tissue-engineered ventricles. Regen Med 2007.2: 391–406.
12. Tanaka Y, Sato K, Shimizu T, Yamato M, Okano T, Kitamori T. A micro-spherical heart pump powered by cultured cardiomyocytes. Lab Chip 2007.7: 207–212.
13. Patel NM, Mohamed MA, Yazdi IK, Tasciotti E, Birla RK. The design and fabrication of a three-dimensional bioengineered
open ventricle. J Biomed Mater Res B Appl Biomater, 2017.105: 2206–2217.
14. Aranaz I, Mengibar M, Harris R, et al. Functional characterization of chitin and chitosan. Current Chemical Biology 20093:203–230.
15. Patel NM, Yazdi IK, Tasciotti E, Birla RK. Optimizing cell seeding and retention in a three-dimensional bioengineered cardiac
ventricle: The two-stage cellularization model. Biotechnol Bioeng 2016.113: 2275–2285.
16. Patel NM, Birla RK. Pulsatile flow conditioning of three-dimensional bioengineered cardiac
ventricle. Biofabrication 2016.9: 015003.
17. Huang YC, Khait L, Birla RK. Contractile three-dimensional bioengineered
heart muscle for myocardial regeneration. J Biomed Mater Res A 2007.80: 719–731.
18. Tao ZW, Mohamed M, Hogan M, Gutierrez L, Birla RK. Optimizing a spontaneously contracting heart tissue patch with rat neonatal cardiac
cells on fibrin gel. J Tissue Eng Regen Med 2017.11: 153–163.
19. Radisic M, Euloth M, Yang L, Langer R, Freed LE, Vunjak-Novakovic G. High-density seeding of myocyte cells for cardiac
tissue engineering. Biotechnol Bioeng 2003.82: 403–414.
20. Swanson M, Clark RE. Dimensions and geometric relationships of the human aortic valve as a function of pressure. Circ Res 1974.35: 871–882.
21. Krogmann ON, Rammos S, Jakob M, Corin WJ, Hess OM, Bourgeois M. Left ventricular diastolic dysfunction late after coarctation repair in childhood: Influence of left ventricular hypertrophy. J Am Coll Cardiol 1993.21: 1454–1460.
22. Salazar BH, Reddy AK, Zewei Tao, Madala S, Birla RK. 32-channel system to measure the electrophysiological properties of bioengineered cardiac
muscle. IEEE Trans Biomed Eng 2015.62: 1614–1622.
23. Ishiwata T, Nakazawa M, Pu WT, Tevosian SG, Izumo S. Developmental changes in ventricular diastolic function correlate with changes in ventricular myoarchitecture in normal mouse embryos. Circ Res 2003.93: 857–865.
24. Choi YH, Stamm C, Hammer PE, et al. Cardiac
conduction through engineered tissue. Am J Pathol 2006.169: 72–85.
25. Beinfield WH, Lehr D. QRS-T variations in the rat electrocardiogram. Am J Physiol 1968.214: 197–204.