Fabrication and characterization of a thick, viable bi-layered stem cell-derived surrogate for future myocardial tissue regeneration

hiPSCs were reprogrammed from human cardiac FBs, and subsequently differentiated into iCMs, as previously reported [4, 47, 48]. Generally, spontaneous contractions were observed in the iCM cultures between days 7 and 10 after the differentiation protocol commenced, with beating numbers increasing up to day 12. Metabolic purification of iCMs via glucose deprivation (RPMI 1640 without glucose, supplemented with sodium DL-lactate and B27⁺, Gibco) for 3–6 d, initiated at day 11, allowed for a population of iCMs that was at least 95% cTnT positive. hiPSCs were maintained at optimal conditions, as previously described [49], on six-well plates coated with Matrigel (Corning), using mTeSR 1 maintenance media (STEMCell Technologies, Canada). hiPSC-ECs (iECs) were differentiated as described previously [21, 50, 51]. Briefly, undifferentiated hiPSCs were seeded into a 0.5 mL fibrin scaffold and treated with CHIR99021 and U46619 in EBM2 medium (Lonza, USA) supplemented with B27^- for 24 h. The medium was replaced with EBM2 supplemented with B27^-, vascular endothelial growth factor (VEGF), erythropoietin, and transforming growth factor β1, and then cultured for 96 h with a media change halfway through. Finally, the scaffold was released and cultured in EGM2-MV medium (Lonza, USA) supplemented with B27⁺, VEGF, SB-431542, with media changes every two days. hiPSC-ECs were purified and enriched by collecting cells positive for CD31 using fluorescence-activated cell sorting device (FACSAria II). Antibodies used for selection along with dilutions are listed supplemental table 1. See supplemental figure 1 (stacks.iop.org/BMM/16/035007/mmedia) for cell characterization and supplemental figure 2 for the proliferation assay performed in to prove the lack of tumorigenic properties of the iCMs utilized.

The supporting fibrin matrix (per milliliter) used for each alternating layer of the cardiac tissue fabrication consisted of the following components as was defined previously [21]: 0.12 mL fibrinogen (25 mg mL⁻¹, Sigma-Aldrich, CAS#9001–32-5), 0.02 mL Matrigel (Corning, #356 235), 0.56 mL of HEPES (20 mM, pH 7.4, Corning), 0.001 mL CaCl₂ (2 M), 0.3 mL DMEM (Gibco, High glucose, #11 965-118), 0.006 mL thrombin (80 U mL⁻¹, MP Biomedicals).

PDMS stilts were fabricated by mixing PDMS (Dow Corning Sylgard 184 Silicone, Product #2 065 622) in a 10:1 elastomer:curing agent ratio and poured into a 100 mm diameter Pyrex Petri dish (Corning, #3 160 102). These were cured at 75 °C for 2 h in an oven, after which custom stilts of 10 mm × 5 mm (2.5 mm thick) were cut. All PDMS structures were autoclaved prior to use in tissue fabrication.

Following cell differentiation, cardiac tissue fabrication commenced (figure 1). Petri dishes (BioLite Cell Culture Treated Dishes, Thermo Scientific) were coated with a 5% pluronic F-68 solution (Gibco, #24 040 032) and incubated at 4 °C overnight. The pluronic solution was removed, and a sterile polycarbonate frame (internal area: 1 × 2 cm²) was attached with a 2% agarose solution. Note that the frames were modified with channels to allow for maximal media contact following tissue fabrication. iCMs were dissociated (STEMdiff Cardiomyocyte Dissociation Medium, STEMCell) and mixed with the fibrin matrix at a concentration of 10 × 10⁶ cells mL⁻¹. Note that the deposition of the iCM layer denotes 'D0' for the remainder of the fabrication process. 400 µl of this solution was quickly deposited into each mold to produce the first layer. Following complete polymerization, the culture medium was added (STEMdiff Cardiomyocyte Support Medium, 2 mg mL⁻¹ ε-aminocaproic acid) and incubated at 37 °C (5% CO₂) for 2 d. The next layer, comprised of iECs, was made in a similar fashion, with the following exceptions: iECs were dissociated using trypsin (0.25% trypsin, 0.1% EDTA, Corning, #25 053 CI) and then mixed with the fibrin matrix at a concentration of 10 × 10⁶ cells mL⁻¹. 200 µl of this solution was quickly deposited into each mold, yielding a 2:1 ratio of iCMs:iECs [21]. Fresh culture medium (10% fetal bovine serum, 2% B27⁺, and 2 mg mL⁻¹ ε-aminocaproic acid, 10 µM ROCK inhibitor in DMEM) was added, following layer polymerization. After 24 h, the frame containing the cardiac tissue surrogate was lifted off of the dish surface and placed on top of custom-cut PDMS stilts, allowing for the tissue to be fully suspended in fresh culture media (2% fetal bovine serum, 2% B27⁺, and 2 mg mL⁻¹ ε-aminocaproic acid in DMEM). Incubation continued for the desired period of time (1–4 weeks in total), with media replacement once per week.

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Samples were fixed in 4% formaldehyde (Pierce, Thermo Scientific, #28 906) for 1 h prior to embedding in either optimal cutting temperature compound (OCT compound, Fisher Health Care, USA) or paraffin for histological analysis (10 µm sections). Whole-mount samples were stored in PBS until staining.

Deparaffinized and rehydrated sections were stained in hematoxylin solution (Mayer's, Merck, 3 min) then working eosin Y solution (2 min). Following dehydration sampled were mounted in Permount and imaged with a bright field microscope (Olympus IX83 epifluorescent microscope).

OCT embedded sections were blocked and permeabilized for 30 min in 10% donkey serum, 3% BSA, and 0.05% Triton-X. Primary antibodies were incubated for 1 h at room temperature (supplementary table 1). Following PBS wash (3 × 5 min), secondary antibodies labeled with fluorescent tags and 4', 6-diamidino-2-phenylindole (DAPI, 100 ng mL⁻¹) were added for 1 h at room temperature (supplementary table 1). Sections were mounted in VECTASHIELD Antifade Mounting Medium and visualized by confocal laser scanning (Olympus FV3000 confocal microscope).

Fixed whole-mount samples were blocked and permeabilized in 10% donkey serum, 10% Tween-20, 3% BSA, 0.05% Triton-X, and 0.2% sodium azide in PBS overnight at 4 °C. Primary antibodies were incubated overnight at 4 °C (supplementary table 1). Samples were washed in PBST (3 × 10 min) then fluorescently labeled secondary antibodies were added overnight at 4 °C with the addition of DAPI. Following washing, tissue was cleared according to the previously described protocol [52] with 3 h incubation in Ce3D clearing agent. Whole-mount constructs were transferred to an Ibidi µ-Slide (#80 286), covered with VECTASHIELD Antifade Mounting Medium, and visualized by confocal laser scanning.

All sample specimens, including PDMS, fresh murine heart tissue and cardiac tissue surrogates were treated with the same procedure described here. Specimen size measurements (surface area, A) used in the determination of the stress from the forces measured (σ = F/A) were determined with a digital caliper (Mitutoyo 500–196-30, Digimatic). Stress relaxation responses, stress (σ) vs. time (t), were characterized using a generalized Maxwell model [53] of order 4 (figure 2(c)) in MATLAB (MathWorks, R2018, see supplemental information for the code). All specimens were transported from the incubator or sacrificed animal in chilled PBS and tested immediately. Stress relaxation testing in compression [54] was done using a Low Force Testbench (TA Instruments, New Castle, DE), where a 10% strain () was applied and held for 4 min on each sample. Samples were strained () at 10% by varying the applied deformation with thickness. Force (F) was measured with a 250 g load cell (figures 2(a), (b)).

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The generalized Maxwell model was defined using the following equations, where E_i and η_i represent the modulus and viscosity of the i^th elements, and σ the measured stress:

$$\begin{align} {\textbf Model:}{\text{ }}{\mathbf \sigma} \left( {\textbf t} \right) =\,\, & {\mathbf \varepsilon} {\mathbf{{ E}_0}} + {\mathbf \varepsilon} {\mathbf{{ E}_1}}\left( {{{\textbf e}^{ - \frac{{{\textbf E_1}}}{{\mathbf{\eta _1}}}t}}} \right) + {\mathbf \varepsilon} {{\textbf E}_2}\left( {{\textbf {e}^{ - \frac{{\mathbf{E_2}}}{{\mathbf{\eta _2}}}t}}} \right) \nonumber \\ & + {\mathbf \varepsilon} {{\textbf E}_3}\left( {\mathbf{e^{ - \frac{{{E_3}}}{{{\eta _3}}}t}}} \right) + {\text{ }}{\mathbf \varepsilon} {{\textbf E}_4}\left( {{\mathbf{e}^{ - \frac{{\mathbf{E_4}}}{{{\mathbf\eta _4}}}{\textbf t}}}} \right)\end{align} \tag{ 1 }$$

For each curve fit, the instantaneous modulus could be determined by the summation of each elastic modulus term (t = 0), whereas the equilibrium modulus was determined at t→∞ where E₀ = σ/. Viscosity was determined as the summation of each viscosity element (η_i ).

$$\begin{align}\begin{array}{l} {\mathbf{Instantaneous}}{\mkern 1mu} {\mathbf{Youn}}{{\mathbf{g}}^{\prime} }{\mathbf{s}}{\mkern 1mu} {\mathbf{modulus}}{\mkern 1mu} \;\left( {{\mathbf{Pa}}} \right) = {{\mathbf{E}}_0} + {{\mathbf{E}}_1} + \\ \qquad \qquad {{\mathbf{E}}_2} + {{\mathbf{E}}_3} + {{\mathbf{E}}_4} \end{array}\end{align} \tag{ 2 }$$

$$\begin{equation}{\mathbf{Equilibrium}}{\mkern 1mu} {\mathbf{Youn}}{{\mathbf{g}}^{\prime} }{\mathbf{s}}{\mkern 1mu} {\mathbf{modulus}}{\mkern 1mu} \;\left( {{\mathbf{Pa}}} \right){\mkern 1mu} = {\mkern 1mu} {{\mathbf{E}}_0}\end{equation} \tag{ 3 }$$

$$\begin{equation}{\textbf Viscosity}{\text{ }}\left( {\mathbf{Pa\,.\,{\textrm s}}} \right) = {\mathbf{\eta _1}} + {\mathbf{\eta _2}} + {\mathbf{\eta _3}} + {\mathbf{\eta _4}}\end{equation} \tag{ 4 }$$

Prior to performing tests on the engineered tissue samples, the validity of this method was tested using PDMS of different crosslinking densities. Moduli and viscosity obtained from these validation experiments were used to confirm the goodness-of-fit for the Maxwell model of order 4 (figure 2(d)). The generalized Maxwell model of order 4 allowed for a good fit to the experimentally obtained data (R² > 0.98)

Studies were performed using healthy murine hearts as a control sample in viscoelastic testing for cardiac tissue surrogates that were generated during this study. All animal procedures were performed in accordance with the guidelines for animal experimentation set forth and approved by Institutional Animal Care and Use Committee (IACUC, APN 20216), School of Engineering, University of Alabama at Birmingham; and conformed to the Guidelines for the Care and Use of Laboratory Animals published by the US National Institutes of Health (2011). Control tissue samples from SCID mouse LVs were obtained after perfusion with saline and arrested in diastole by injection with 25 mM potassium chloride. Animals were gender-randomized.

Ultrastructural analysis of the thick cardiac tissue surrogates was done via TEM imaging. After culturing for the required amount of time (1–4 weeks), the tissue surrogates were fixed in a 2.5% glutaraldehyde solution for 1 h at 4 °C prior to being delivered to the UAB High-Resolution Imaging Facility for further processing. Sample blocks were sectioned using a diamond knife for clean and even sections. Samples were mounted and viewed using a Tecnai Spirit T12 TEM. Images were collected for analysis at each time point of interest (1 week, 2 weeks, and 4 weeks of tissue culture).

To assess intercellular coupling in the tissue surrogates, we used systems of very small and closely spaced electrical sensors custom-fabricated into linear arrays that are integrated with adjacent instrumentation to provide low-amplitude stimulation current to the interstitial compartment of 15 preparations. Within each array, alternating current stimuli were delivered between the outer pair of electrodes at frequencies of 10 Hz to 4000 Hz as described previously [21, 55, 56]. Each stimulus established a 3D interstitial potential field that was then sensed as a voltage difference between sensors that formed an inner pair of electrodes that in turn allowed the identification of a microscopic composite impedance that included both real (tissue resistivity) and imaginary (tissue reactivity) components. At relatively low frequencies, supplied current remains primarily interstitial, and as a consequence, the current-to-voltage ratio provides information regarding the strength of interstitial electrical coupling. As frequency increases, supplied current shunts to the intracellular compartment via membrane capacitance and intracellular coupling strength becomes essential in assessing the response. Tissue resistivity and reactivity were identified during four sequential acquisition intervals in each preparation.

To assess action potential duration, conduction velocity, and minimum pacing cycle length, LbL tissue surrogates were stained with a voltage-sensitive dye RH-237 (2.5 μM) for 10 min, transferred to a perfusion chamber mounted on an inverted microscope (scope details). Samples were perfused with Hank's balanced salt solution (HBSS) at approximately 37 °C. Sample pacing/stimulation was done with a bipolar electrode consisting of a glass pipette filled with HBSS and a silver wire coiled around its tip. The electrode tip was positioned at the sample's edge using a micromanipulator. Rectangular stimulation pulses were used, with a duration of 2 ms and current strength 1.5 times the excitation threshold. Fluorescence was excited with a 200-W Hg/Xe arc lamp and recorded with a 16 × 16 photodiode array (Hamamatsu) at a spatial resolution of 110 μm per diode as previously described [57]. Excitation light was filtered at 532–587 nm, and emitted fluorescent light was filtered at >650 nm. To eliminate motion artifacts caused by the sample's spontaneous contractions, the perfusion solution was supplemented with 5 μM of blebbistatin. Isochronal maps of the activation spread were constructed from activation times measured at 50% of the maximum action potential amplitude. Conduction velocity was calculated at each recording site from local activation times and averaged across the whole mapping area. Action potential duration was measured at 50% and 80% of signal recovery (APD₅₀) and (APD₈₀), respectively.

All image quantification analyses were performed with ImageJ. Where indicated, arbitrary units are representative of a pixel count and intensity for each sample.

For statistical analysis, data are shown in the form mean ± SEM. Significance was chosen as p < 0.05. This was determined using either Student's t-test or ANOVA where applicable. One sample Wilcoxon tests (α = 0.05) were used in the viscoelastic statistical analyses of between cardiac tissue surrogates (n = 3) and native murine LV tissue (n = 5). These analyses were performed utilizing GraphPad Prism8 data analysis software package.

Tissue surrogates produced with the optimized method described in figure 1 yielded structures of 1.73 ± 0.075 mm in thickness after 1 week in culture (figure 3(a)). Following the first 24 h post-iCM layer deposition (on D1), the beating-rate (per min) of each tissue surrogate was determined (figure 3(b)). This yielded a rate of 74 ± 8.4 beats min⁻¹ (bpm). On the second day, the beat-rate lowered significantly (p = 0.01) to 63 ± 5.8 bpm. After iEC layer addition on D3 and an additional 3 d of culture (D6), the beating-rate on the sixth day had decreased further to 29 ± 4.1 bpm (p = 0.0008). Whether this significant decrease in bpm is due to paracrine signaling from the iECs or whether the internal cardiac pacemaker system become somewhat quiescent, remains to be seen.

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Alterations in ECM composition between weeks 1 and 4 were observed via H&E staining (figures 3(c)–(h)). Histology suggests increases in structural ECM components such as collagen (increased intensity as well as the distribution of eosinophilic staining, figures 3(c)–(h). With increased culture time, compaction of the structures became more apparent (figure 3(f)), and structures, although initially somewhat disorganized, resembled in vivo muscle. Structure coupling was confirmed by not only the synchronous macroscopic beating of the tissue surrogates (figure 3(b)) but also through the fusion of the two layers as visualized on H&E sections. Another observation was the iEC migration out of their originally deposited layer into the resident iCM layer as is clear from the somewhat acellular appearance of the top layer of the tissue surrogate (figures 3(c)–(h)), Cell migration from the iEC later into the iCM resulted in a localized elevated concentration of iECs near the layer-fusion site, and was primarily evident during the initial culture stages (after 1 week, figures 3(c)–(e)) as opposed to later culture stages (week 4, figures 3(f)–(h)). Due to the difficulty in distinguishing between the different cells and their distribution throughout the structure based on H&E alone, further specific immunofluorescent staining was completed and quantified (figure 5). The effect of this migration on the viability of the cardiac surrogates was determined via immunofluorescent staining and subsequently quantified. It should be noted that structures visualized here via H&E staining appear thinner than those preserved in OCT. This difference in overall tissue thickness can be attributed to the multiple dehydration-related steps required during the H&E processing of the hydrogel structures leading to decreased preservation of the original architecture.

The viability of the tissue was determined with pMLKL staining (figure 4), which specifically stains for cell necroptosis [58] and is associated with inflammatory markers [59]. The pMLKL-positive cells were quantified as a percentage (%) of the total number of cells present (i.e. normalized to the DAPI staining). No significant differences were found in the degree of necrotic staining for any of the time points analyzed. The necrotic staining was quantified (figure 4(c)) as 4.5 ± 2.3% after 1 week, 4.3 ± 1.1% after 2 weeks, and 5.6 ± 1.4% after 4 weeks in culture (see supplemental figure 3 for staining at week 4). These values were comparable with the ∼4% pMLKL positive staining obtained by Gao et al previously for their large tri-lineage iPSC-derived thick (1.25 mm) cardiac muscle patches after 1 week in culture [21].

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Imaging of the fluorescently stained sections (figure 5) showed iEC migration into the iCM region of the dual-layered tissue surrogates. Evidence of migration was initially seen after day 5 and continued throughout the culture period. This migration process continued over the 4 weeks of culture, with clear integration into the iCM layer observed from the 2D sections by the 4th week (figure 5(a)). In order to confirm vasculature formation as well as establish a timeline for the formation process, whole-mount staining was employed. Staining of samples after 2 weeks of culture (figure 5(c)) showed branching vascular networks. Whole-mount staining further confirmed the observations regarding iEC integration and elucidated that iCM orientation was random and did not yet present with the degree of cell alignment seen in native myocardial tissue (figure 5(c)). Clear delineation of cell distribution can be seen in supplemental figure 4. Quantification of each cell type at weeks 1, 2 and 4, respectively, showed that the amount of CD31 expression relative to cTnT expression remained stable over the 4 week period with no statistically significant differences between the expression ratios at each time point (n = 3, supplemental figure 4(c)). Furthermore, no statistically significant differences between the cTnT expression relative to the DAPI expression was noted either (n = 3, supplemental figure 4(d)).

Engineered tissue modeling over the 4 weeks yielded excellent fits to the experimentally obtained results, as R² > 0.7 for all samples from weeks 1 and 2 respectively, while R² > 0.88 for samples from week 4 (figure 6(a)). Instantaneous moduli of 4.72 ± 0.84 kPa, 5.82 ± 0.47 kPa, and 4.85 ± 0.16 kPa were measured for week 1, week 2, and week 4 samples, respectively (figure 6(b)). The standard deviation in the samples decreased as the culture time increased from 1 week to 2 weeks to 4 weeks. Viscosity modeling yielded values of 8.03 × 10⁴ ± 2.42x10³ Pa•s, 3.35 × 10⁴ ± 6.31x10³ Pa•s, and 6.70 × 10⁴ ± 2.14x10⁴ Pa•s respectively. Tissue surrogate viscoelastic properties (n = 3 for each group) were compared to those of healthy LV tissue from SCID mice (n= 5). Mouse LV tissue instantaneous modulus was measured at 6.43 ± 1.13 kPa, while viscosity was 2.00 × 10⁵ ± 8.32x10⁴ Pa•s. Equilibrium moduli of 1.08 ± 0.31 kPa, 1.31 ± 0.37 kPa, and 0.40 ± 0.10 kPa were obtained for week 1, week 2, and week 4 samples, respectively (figure 6(c)). The equilibrium modulus for the control mouse LV samples was observed at 2.81 ± 0.48 kPa. Non-parametric tests (one sample Wilcoxon) showed that there were no statistically significant differences between the viscoelastic properties of the cardiac tissue surrogate and the healthy murine LV tissue (α = 0.05), even with the differences in properties noted for the tissue surrogates at different time points.

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During the 4-week culturing period, the ECM underwent multiple phases of remodeling. The degree to which this very dynamic process occurs was monitored and quantified in vitro through analysis of individual matrix component degradation and deposition. For ECM analyses, three samples were analyzed at each time point, with at least five images per matrix component. Since the fibrin used in the matrix fabrication consisted of both bovine thrombin and fibrinogen, an antibody targeting a bovine specific epitope was used in its analysis. An antibody with a fibrinogen β-chain that matched the β-chain section of the fibrinogen used to fabricate the fibrin matrix was used to stain and analyze the evolution of fibrin over 4 weeks in culture (figure 7(a)). Analysis showed discernable degradation of the fibrin matrix over this period. Since the engineered tissue remained structurally intact and previous research has shown that ECM remodeling occurs during iCM culturing [60, 61], samples were analyzed for additional ECM components, including collagen 1 (Col1), collagen 3 (Col3), collagen 4 (Col4), laminin (Lam) and fibronectin (FN) (figure 7(b)). Quantification showed a nearly three-fold increase in collagen one production from weeks 1–4. A general trend was observed in basal membrane component production, i.e. the relative amounts of both collagen four, as well as laminin, decreased from week 1 to week 2, to week 4. The variations in structural ECM components, like collagen 1, affect the overall stiffness of the tissue surrogates, whereas dense basement membrane components, including laminin and collagen 4, play more of a role in viscous properties [62–64]. The decreasing viscosity of the tissue surrogates (figure 6(b)) is likely linked to the reduction in basement membrane ECM components. Furthermore, week 2 saw a notable increase for fibronectin, which is highly expressed in the heart during early stages of embryogenesis and has also been shown to be vital in the vasculogenesis process [65]. The enhanced expression of fibronectin noted after 2 weeks in culture may be linked to the formation of the branched vessel-like structures noted in figure 5(c). In addition to altering viscosity properties of the structure, collagen 4 likely has another functional role, specifically that of an anti-angiogenic signaling cue [66, 67]. The statistically significant increase in collagen 4 in week 2 coincides with an even more notable statistically significant increase in fibronectin expression (figure 7(b)). This phenomenon can most likely be attributed to an 'on-off' switch regulating the vessel formation process occurring in the tissue surrogate.

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Alterations noted in the ECM via H&E and immuno-fluorescent staining (figures 3(c)–(h)) were also confirmed via TEM (figure 8). As culture-time increased from 1 to 4 weeks (figures 8(a) vs (c) vs (e), there was an increase in the number of gap junctions (GJ) observed between the cells (figure 8(d)). Increases in the number of GJ (see supplemental figure 5 for immunofluorescent staining of connexin 43 as well as its quantification), along with the statistically significant increase in sarcomere length (distance between z-bands) from 1.53 ± 0.02–1.68 ± 0.12 μm between weeks 1 and 2, with a further increase to 1.70 ± 0.02 μm at week 4 (figure 8(g)) suggest that the fabrication method along with the extended culturing promote a degree of functional maturation [68]. These sarcomere lengths are physiologically relevant and compare with those observed in both neonatal rats and sheep (1.5–1.9 µm) [69, 70].

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Functional performance was assessed by t-tubule formation, intracellular GJ formation (4EMS) as well as optical mapping to determine conduction velocities through the tissue. All mapping of surrogates was performed after 1 week in culture. Staining and subsequent quantification of the tissue surrogates for both JP2 (figure 9(a), supplemental figure 6) and RyR (figure 9(b), supplemental figure 6) expression suggested that there was a significant increase in the number of t-tubules formed between the first and fourth week of tissue culture. Resistivity–reactivity spectra (figure 9(c)) compared favorably with spectra from our earlier report with rabbit LV myocardium [55], supporting GJ intracellular communication developed in these cardiac tissue surrogates to an extent similar to that in native myocardium. The heatmap generated from optical mapping (figure 9(d)) shows a lack of arrhythmogenicity in the bi-layered structures. Furthermore, structures showed conduction velocities of 16.9 ± 2.3 cm s⁻¹, with the ability to be paced to 300 ms (3.33 Hz) from 800 ms (1.25 Hz), before being recovered to 800 ms pacing. ADP₅₀ and ADP₈₀ were noted as 128 ± 12.0 and 186 ± 17.7 ms, respectively.

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