Cardiovascular disorders are prevalent in today’s society and adversely affect the lives of millions of patients around the globe.1,2 This epidemic is progressively worsening with an aging population and an increase in obesity.3 More than 50% of patients with heart disease do not respond to current pharmacological therapies4; to date, heart transplantation is the most effective option for patients suffering from end-stage heart failure. However, in addition to the logistical and medical complexity of this therapeutic option, donor organ shortage represents a major limitation, and up to 20% of the patients die while on the waiting list.5 Tissue engineering strategies have evolved in response to this need and are focused on the fabrication of artificial tissue constructs in a controlled laboratory environment.6–10
Research in the area of heart muscle tissue engineering is focused on the development of functional three-dimensional artificial heart muscle (3D-AHM), and more recently, organ fabrication technologies have been developed to support the fabrication of complete artificial hearts.11–13 The most common strategy to fabricate 3D-AHM has been to culture neonatal cardiac myocytes within a suitable 3D scaffold, resulting in artificial tissue that replicates a partial subset of mammalian heart muscle functions. There is a vast body of literature describing the fabrication of 3D-AHM, including publications from our group.14–19 These studies have often been coupled with endeavors designed to evaluate the role of physiologic stimulation (mechanical stretch, electrical stimulation, and perfusion) to guide 3D tissue formation and function. These constructs are intended to be used clinically to patch areas of myocardial infarction. In addition, 3D-AHM models provide a platform for detailed study of critical events in organogenesis, such as establishment of cell–cell communication, cytoarchitecture, and extracellular matrix (ECM) formations. The field of tissue engineering has matured from initial proof of concept studies to the point of demonstrating feasibility of artificial organs, with recent examples of whole heart,20–22 lung,23 and kidney models.24 Furthermore, tissue engineered tracheas25,26 and bladders27 have been used clinically in patients.
Through culture of neonatal cardiac cells on fibrin gel, our group has fabricated 3D contractile cardiac tissue15,16 and subsequently attached 3D cardiac tissues to acellular blood vessels and heart scaffolds to support bioartificial pump and heart fabrication, respectively.17,18 We designed the current framework to fabricate a bioartificial heart (BAH). The model proposed herein is based on acellular scaffolds coupled with intramural injection recellularization strategies.
All animal protocols were approved by the Institutional Animal Care and Use Committee (IACUC) at the University of Houston, in accordance with the “Guide for the Care and Use of Laboratory Animals” (NIH publication 86-23, revised 1996). All materials were purchased from Sigma-Aldrich (St. Louis, MO) unless otherwise specified. Figure 1 shows the schematic method for fabrication of BAHs.
Isolation of Primary Cardiac Cells
Cardiac cells were isolated from the hearts of 2–3 day old neonatal Sprague-Dawley rats using an established method.19 Each heart was cut into 3–4 pieces in an ice-cold phosphate buffer saline (PBS) consisting of 116 mM NaCl, 20 mM HEPES, 1 mM Na2HPO4, 5.5 mM glucose, 5.4 mM KCl, and 0.8 mM MgSO4. After blood cells were rinsed out, heart pieces were transferred to a dissociation solution consisting of 0.32 mg/ml collagenase type 2-filtered (Worthington Biochemical Corporation, Lakewood, NJ) and 0.6 mg/ml pancreatin in previously mentioned PBS. The hearts were cut into 1 mm3 pieces and then transferred to an orbital shaker and maintained at 37°C for 30 minutes at 60 rpm. At the end of the digestion process, the supernatant was collected in 3 ml of horse serum to neutralize the enzyme and centrifuged at 1000 rpm for 5 minutes at 4°C. The cell pellet was resuspended in 5 ml horse serum and kept in an incubator at 37°C supplied with 5% CO2. Fresh dissociation solution was added to the partially digested tissue, and the digestion process was repeated an additional 2–3 times. Cells from all the digests were pooled, centrifuged, and suspended in culture medium consisting of M199 (Life Technologies, Grand Island, NY), with 20% F12k (Life Technologies), 10% fetal bovine serum, 5% horse serum, 1% antibiotic-antimycotic, 40 ng/ml hydrocortisone, and 100 ng/ml insulin. Cell viability was analyzed by trypan blue (4%) staining according to the manufacturer’s protocol.
Preparation of Acellular Scaffolds
We used a detergent based decellularization process that has been developed and used extensively for several cardiovascular and musculoskeletal tissue engineering applications.18,28–30 Hearts were obtained from 3 to 4 month old female Sprague-Dawley rats and immediately washed in a PBS solution to prevent blood coagulation. The hearts were then sequentially incubated in 15 ml of a series of decellularization solutions. Five solutions were made with the following compositions: solution 1 (S1), 80% glycerol + 0.9% NaCl + 0.05% NaN3 + 25mM ethylenediaminetetraacetic acid (EDTA); S2, 4.2% sodium deoxycholate + 0.05% NaN3; S3, 1% sodium dodecyl sulfate (SDS) + 0.05% NaN3; S4, 3% Triton X-100 + 0.05% NaN3; and S5, 0.05% NaN3. Over a 16 day period, the hearts were exposed to the solutions in the following order: days 1 and 2 in S1; days 3–5 in S2; days 6–7 in S1; days 8–9 in S3; days 10–11 in S4; days 12–13 in S3; and days 14–16 and after in S5. Figure 1A–D shows the changes of the cadaveric heart in the decellularization process. Before use, the acellular heart was cannulated to the aorta using a 16 gauge needle and perfused and rinsed with sterile PBS for 15 minutes.
For assessment of the vascular network, 0.1% trypan blue PBS was perfused within the heart scaffold to examine the dilation of the blood vessels (Figure 1E). For assessment of the decellularization quality, the whole decellularized heart scaffold was sectioned, and sections were stained with Masson’s trichrome reagents and also examined with immunofluorescence staining of α-actinin, cardiac troponin I (cTnI), collagen type I, and nuclei (Figure 1F–J) to see whether there were any cardiac muscle or nuclei left.
Fabrication of Bioartificial Hearts
The acellular heart was suspended within a culture vessel and placed in a Langendorff perfusion system, and perfused with culture media previously mentioned. Culture media reservoir was maintained at an elevated position (100 cm) to drive the gravity fed perfusion. The system was housed in an incubator for temperature and pH regulation (Figure 2A and B). The 2 ml cell suspension containing 50 million cells was injected into multiple regions in the acellular heart. After a 24 hour culture period, a second set of 50 million cells was injected into the same heart. The cells in the effluent media within the first 2 hours were pulled; cell viability was analyzed as described previously, and then, these cells were injected to the heart again. Perfused media was collected and recirculated to the heart using a peristaltic pump, which was placed outside the incubator; fresh media was replaced every 2 days.
Electrocardiogram and Ventricular Pressure Measurements
Electrocardiogram (ECG) of the heart was measured using Octal Bio Amp (ML138, ADInstrument, Colorado Springs, CO). Data were acquired through a 16 channel PowerLab system (PL3516/P, ADInstruments). Electrocardiogram was detected by inserting the needle anode (MLA1213, ADInstruments) into the left ventricle (LV) anterior wall, attaching the cathode onto the right atrial appendage and the ground onto the aortic wall. LabChart (ADInstruments) was used for data analysis. Bioartificial heart ventricular pressure was measured using a 3.5 F high fidelity, catheter-tipped micromanometer (model SPR-524, Millar Inc., Houston, TX). The left atrium was cut, and the catheter was inserted through the mitral valve and advanced into the LV chamber. The output from the Millar catheter was recorded using PowerLab system (Figure 2C).
For comparison of the ECG signal and LV pressure with that from normal adult hearts, 3–4 month old female Sprague-Dawley rats were anesthetized with ketamine (10 mg/kg) and xylazine (20 mg/kg). Hearts were removed and immediately mounted onto Langendorff perfusion system and perfused with the same media as described previously.
Electrical Pacing and Isoprenaline
Electrical pacing was performed 4 days after the second cell injection by attaching the stimulating electrodes (MLA0320, ADInstruments) on the anterior wall of the heart (Figure 2R) at frequencies of 0.25 and 1–6 Hz (in 1 Hz increments; Figure 2D). Isoprenaline at 1, 10, and 100 nM were added sequentially to the perfusion on day 4, and the changes of heart rate (HR) and LV pressure were recorded through a PowerLab system.
Histology and Immunofluorescence
On day 8 after the second cell injection, the recellularized heart was rinsed with PBS for 15 minutes and then removed from the perfusion system. The heart was dried with a paper towel, weighed, and placed in a peel-away disposable embedding mold (VWR International, Radnor, PA), frozen in liquid N2, and then immediately immersed in Tissue Tek OCT compound (VWR International, Radnor, PA) and placed in a –80°C freezer. Once the OCT compound solidified, each sample was sliced using a cryostat (Thermo Fisher Scientific, Waltham, MA). Whole heart longitudinal sections from the base to apex of the heart were cut at a thickness of 10 μm. The sections were placed on VWR Microslides for preparation of morphological and immunofluorescence examinations. For assessment of cellular distribution within BAHs, whole heart sections were stained with Masson’s trichrome reagents according to the manufacturer’s protocol. Section images were taken under a light microscope, and the images of whole heart sections were stitched together using the Grab Large Image feature within NIS Elements 4.13 (Nikon Instruments Inc., Melville, NY) software.
Whole heart sections of 10 μm thickness were fixed in ice-cold acetone for 10 minutes; nonspecific epitope antigens were blocked with 10% goat serum at room temperature for 1 hour. Sections were incubated with specific mouse anti-α-actinin antibody (Sigma, A7811) 1:200, rabbit anti-cTnI antibody (Abcam, ab47003) 1:100, mouse anti-vimentin (Sigma, V6630) 1:40, rabbit anti-collagen type I (Abcam, ab34710) 1:100, rabbit anti-connexin 43 (Cx43; Abcam, ab11370) 1:100, rabbit anti-von Willebrand factor (vWF; Abcam, ab6994, Cambridge, MA) 1:750, and rabbit anti-ki67 (Abcam, ab15580) 1:100 at room temperature for 1 hour. Subsequently, sections were treated with goat anti-mouse or goat anti-rabbit secondary antibodies (Alexa Fluor 488 and Alexa Fluor 546, Life Technologies) 1:400 at room temperature for 1 hour. Nuclei were counterstained with 4,6-diamidino-2-phenylindole (DAPI; 2.5 μg/ml) for 5 minutes at room temperature. Fluorescence images were obtained with a Nikon C2+ confocal laser-scanning microscope (Nikon Instruments Inc.).
Results are presented as mean ± standard deviation (SD). Comparisons among groups were made with a one-way analysis of variance (ANOVA) followed by the Bonferroni post hoc comparison test. In all tests, differences were considered statistically significant at p < 0.05.
Decellularization of Rat Hearts
Based on visual inspection, we observed a gradual decrease in the opacity of the rat hearts through the course of decellularization protocol. Freshly isolated rat hearts were visibly pinkish red (Figure 2A), and during the course of the decellularization process, the color of the hearts changed to a light brown and whitish color (Figure 2B–D). The decellularized heart was retrograde perfused first with PBS and then 0.1% trypan blue PBS to assess vascular integrity. Figure 2E shows representative images of the dilated left descending coronary artery. To validate the removal of cellular components, sections from decellularized hearts were stained with Masson’s trichrome stains and also examined with immunofluorescence staining of α-actinin, cTnI, collagen type I, and nuclei. Figure 2F is a Masson’s trichrome staining image of a longitudinal section of the whole decellularized heart. As shown in Figure 2G, there were no signs of cardiomyocyte specific proteins or nuclei in any whole heart scaffold sections, which were also further verified by immunofluorescence staining of α-actinin (Figure 2H), cTnI (Figure 2H), and DAPI in the whole heart section. The stain for collagen type I is positive (Figure 2J) in the whole heart section, indicating that the natural heart framework remains intact.
Weight of Bioartificial Hearts
Of the cells injected, 43.5 ± 3.7% (n = 4) were washed away within 2 hours after each injection. However, these discharged cells were injected back to the scaffold. Twelve heart scaffolds were recellularized in the incubator. Two were contaminated 2 days after cell injection, and another 2 were terminated during the culture because of power outrage. Eight days after the second cell injection, recellularized hearts were taken off, drained and weighed. Table 1 shows the average weights of these recellularized hearts. Compared with the natural hearts (1.16 ± 0.13 g, n = 6), the recellurized hearts (0.42 ± 0.07g, n = 6) were much lighter (p < 0.01); however, there was no significant difference compared with the acellular hearts (0.41 ± 0.08 g, n = 6; p > 0.05).
Functional Properties of Bioartificial Hearts
Spontaneous contraction was observed on BAHs within 2 days of the second cell injection. Electrocardiogram signals were detected on the LV anterior walls of recellularized hearts 2 days after the second recelluarization. As shown in Figure 3A, the ECG QRS pattern from the LV anterior walls of the recellularized hearts was either QS (QS = 0.98 ± 0.12 mV, n = 3) or RS (R = 0.25 ± 0.10 mV, n = 3); however, the ECG QRS patterns from that of normal hearts were all qR (R = 4.09 ± 0.86 mV, n = 6; Figure 3B).
Figure 4A shows maximum amplitudes of contraction from the present recellularized hearts on day 4 or 5, after which they decreased gradually; on day 8, contraction was hardly visible. Video (see Supplemental Digital Content 1, http://links.lww.com/ASAIO/A64) shows the contraction of a recellularized heart on days 2, 4, 6, and 8. The recellularized hearts were electrically paced on day 4; as shown in Figure 4B, the highest paced response frequency was 5 Hz. Addition of isoprenaline 1, 10, and 100 nM sequentially to the perfusion system caused dramatic increase in HR, and premature beating occurred when isoprenaline was at 100 nM (Figure 4C).
The rate of contraction was determined by ECG and further verified by the visible contraction of the heart. Heart rate on day 2 (118 ± 6 bpm, n = 6) was significantly higher when compared with days 4 (85 ± 5 bpm, n = 6), 6 (68 ± 6 bpm, n = 5) and 8 (55 ± 5 bpm, n = 5; p < 0.01). Compared with days 6 and 8, HR on day 4 was still faster (p < 0.01); HR on day 8 was the slowest when compared with other day points (p < 0.05 or p < 0.01; Figure 4D). Left ventricle pressure was detected on day 2, and the greatest single LV pressure was 0.61 mm Hg, which was measured from a recellularized heart on day 4 after the second cell injection. Compared with days 2 (0.20 ± 0.11 mm Hg, n = 6; p < 0.01), 6 (0.28 ± 0.05 mm Hg, n = 5; p < 0.05), and 8 (0.18 ± 0.04 mm Hg, n = 5; p < 0.01), the average LV pressure on day 4 (0.43 ± 0.09 mm Hg, n = 6) was the greatest; however, it decreased on day 6 and further on day 8 (Figure 4E). The addition of isoprenaline to the perfusion system caused a significant increase of HR on day 4 (p < 0.01) when compared with that of 0 nM isoprenaline (77 ± 7, n = 6; Figure 4C and F). Compared with 1 (119 ± 10 bpm, n = 6) and 10 nM isoprenaline (138 ± 8 bpm, n = 5), HR with 100 nM isoprenaline (160 ± 9, n = 5) was significantly faster (p < 0.01).
Histological Properties of Bioartificial Hearts
We assessed the histological properties of BAHs using Masson’s trichrome and immunofluorescence stainings (Figure 5). Both of these methods confirmed the presence of a large number of cells that were distributed throughout the acellular graft. Figure 5A1–A3 illustrates sequential sections from right ventricle (RV) to LV stained with Masson’s trichrome reagents. We obtained images from different regions of the recellularized hearts both with Masson’s trichrome and immunofluorescence stainings. A higher concentration of cardiac tissue was observed in the LV anterior (Figure 5C1 and C2), inferior (Figure 5D1 and D2), and LV lateral walls (Figure 5E1 and E2) than in the RV anterior wall (Figure 5B1 and B2). The immunofluorescence images from the LV inferior wall were positive for the features of cardiomyocytes including α-actinin (Figure 5F) and cTnI (Figure 5G), which demonstrate myocardiocytes. Positive staining of vimentin (Figure 5I) and collagen type I (Figure 5J) indicated cardiac fibroblasts and potential ECM production and support. Positive stainings for connexin 43 (Figure 4K) illustrated that the recellularized cardiac tissue maintained electromechanical coupling, which sustains electrical propagation between cardiac cells. In these sections, there were also abundant vWF signals (Figure 5L), which are markers of endothelial cells, and ki67 signals (Figure 4M), which denote cellular division.
The field of tissue engineering has progressed to a point where researchers can endeavor to whole organ engineering. Recent advances in this field have broadened the scope from tissue constructs to complex hollow-chamber structures all the way to entire artificial organs.20–22,31–33 The work in this study provides a framework for a BAH fabrication.
The decellularization process in this study is a 16 day period of submerging the cadaveric hearts in five different detergent solutions. Masson’s trichrome and immunofluorescence stainings demonstrated that there were no cardiac cells left in these scaffolds. Ott et al.20 and Yasui et al.22 prepared their acellular rat hearts by perfusion 0.5–0.1% SDS overnight or 12 hours and followed by 1% Triton-X 100 for 30 minutes. The key point is to make sure that the detergent dissolves just enough materials to release all cells but not to alter the overall structure of the heart.
In this study, we used primary cardiac cells isolated from neonatal rat hearts. Although these cells cannot be used for translational work, they represent critical tools during early stage development. There is a vast body of literature demonstrating the feasibility of primary cardiac myocytes in cardiac tissue engineering.32,34–36 From a model development standpoint, primary cardiac cells provide an excellent platform to understand the variables that affect BAH fabrication, formation, and function. Our previous experiments showed cell viability of the current isolation method is 81.0 ± 2.2%. Our data revealed that the properties, such as spontaneous contraction rate, pacing response frequency of a cardiac patch fabricated with neonatal cardiac cells are highly affected by the primary cell plating density and culture time.19 The same phenomenon holds true for the present BAH fabrication. We observed that HR and LV pressure decreased with culture time (Figure 4A, D, and E), which was mainly a result of the overpopulation of the cardiac fibroblasts. Using fluorescence activated cell sorting technique, Banerjee et al.37 showed that the population of cardiomyocytes in the day 1 rat heart was 62%, cardiac fibroblasts comprised 30%, and nonmyocytes and nonfibroblasts consist of 8%. After 1–3 days of culture, a nongenetic method by Nguyen et al.38 demonstrated that the population of cardiomyocytes was 25.8% and nonmyocytes was 54.2%. Thus, by adding 10 μM cytosine arabinoside, a fibroblast proliferation inhibitor, Yasui et al.22 were able to fabricate a BAH with neonatal cardiac cells, which was capable of beating up to 30 days. Aside from primary cells, cardiomyocytes derived from embryonic stem cells (ESC)39 or human induced pluripotent stem cells (iPSC)21 emerged as a new source of cells for BAH fabrication; however, purified cardiac progenitors might result in less teratoma formation.40
The cellularization strategy represents a significant scientific challenge in organ engineering. It is of great significance to control the position and distribution of the delivered cells during 3D culture as well as for organ fabrication. We explored the use of direct cell transplantation, whereby isolated cells are delivered to the acellular scaffold using multiple intramural injections. Although this strategy is not exact and a significant number of cells are still lost because of the cell suspension going into the ventricular chambers or spilling out, direct cell transplantation strategies have proven efficient in establishing organ level functionality. Our cellularization strategy did result in some level of organ function, as demonstrated by ECG and pressure measurements. The ECG QRS patterns from the LV anterior walls of the current BAHs were either QS or Rq, which might result from pacemaker cell location. It was inferred from QRS patterns (qR; Figure 3) measured in the LV anterior walls of normal adult rat hearts that Rq type QRS complexes are produced when pacemaker cells localize to the upper part of the ventricular chamber, and QS type QRS complexes arise from pacemaker localization to lower regions of the ventricle
The greatest single LV pressure in the current study is only 0.61 mm Hg, and the average LV pressure is 0.43 ± 0.09 mm Hg (n = 6) on day 4 (Figure 4E). Ott et al.20 used the same injection method to deliver about 60 million cells into each rat heart scaffold, and 20 minutes later, 46% of the cells were lost. They recorded a 2.5 mm Hg LV pressure. Yasui et al.22 delivered 100 million cardiac cells by infusion to each rat heart scaffold. Although they did not evaluate how many cells were lost after infusion, they achieved a 0.75 mm Hg LV pressure. The current study, through serial injections, delivered a total of 100 million cardiac cells to each heart scaffold, of which 43.3% are washed out 2 hours after injection, which were recollected and reinjected. As shown in Table 1, compared with the heart scaffolds (0.41 ± 0.08 g), the recellularized hearts (0.42 ± 0.07 g) were not significantly heavier. Cadaveric hearts weigh about 1.16 g, 0.74 g more than a recellularized heart. Injection of 100 million cardiac cells only increased the heart weight 0.01 g. The present method only achieved about 36% (0.42 of 1.16 g) of the mass of an adult rat heart. Acquiring sufficient cells for a single rat BAH fabrication is tough.
During normal mammalian function, the heart is exposed to a myriad of complex signals, including electromechanical cues, fluid stresses, and changes in the hormonal and growth factor environment. There are temporal and spatial variations in all of these signals, which, in turn, regulate the phenotype of mammalian hearts during normal physiologic culture. It is fairly obvious that during in vitro BAH culture, we need to recapitulate this complex in vivo culture environment. The current system only provides continuous media perfusion and lacks many other physiologic signals.
The current study described a framework for the fabrication of a BAH that replicated a partial subset of mammalian heart functions. Bioartificial hearts fabricated by freshly isolated neonatal rat cardiac cells will spontaneously begin beating within 2 days after cell seeding, but their contraction rate and LV pressure decrease with culture. The present model only accounts for about 36% of the mass of an adult rat heart. We need to explore methods to increase cellular mass and maturity to better replicate native tissues. For the next generation of BAH fabrication, a mixture of neonatal cardiac cells with ESC-derived or iPSC-derived cardiac progenitors might be the new candidate for a cell source. Additional electromechanical cues, fluid stresses, and changes in the hormonal and growth factor environment may also improve BAH fabrication.
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