The influence of electrically conductive and non-conductive nanocomposite scaffolds on the maturation and excitability of engineered cardiac tissues

Ali Navaei a, Kiarash Rahmani Eliato bc, Robert Ros bcd, Raymond Q. Migrino ef, Brigham C. Willis fg and Mehdi Nikkhah *a
aSchool of Biological and Health Systems Engineering (SBHSE), Arizona State University, Tempe, Arizona 85287, USA. E-mail: mnikkhah@asu.edu; Fax: +480-727-7624; Tel: +480-965-0339
bDepartment of Physics, Arizona State University, Tempe, AZ 85287, USA
cCenter for Biological Physics, Arizona State University, Tempe, AZ 85287, USA
dBiodesign Institute, Arizona State University, Tempe, AZ 85287, USA
ePhoenix Veterans Affairs Health Care System, Phoenix, AZ85012, USA
fUniversity of Arizona College of Medicine, Phoenix, AZ 85004, USA
gPhoenix Children's Hospital, Phoenix, AZ 85016, USA

Received 31st August 2018 , Accepted 23rd October 2018

First published on 24th October 2018


Utilization of electrically conductive nanomaterials for developing nanocomposite scaffolds has been at the center of attention for engineering functional cardiac tissues. The primary motive in the use of conductive nanomaterials has been to develop biomimetic scaffolds to recapitulate the extracellular matrix (ECM) of the native heart and to promote cardiac tissue maturity, excitability and electrical signal propagation. Alternatively, it is well accepted that the inclusion of nanomaterials also alters the stiffness and nano-scale topography of the scaffolds. However, what is missing in the literature is that to what extent the sole presence of nanomaterials within a scaffold, regardless of their conductivity, influences the maturation and excitability of engineered cardiac tissues. To address this knowledge gap, we developed four different classes of gelatin methacrylate (GelMA) hydrogels, with varied concentrations, embedded electrically conductive gold nanorods (GNRs) and non-conductive silica nanomaterials (SNPs), to assess the influence of matrix stiffness and the presence of nanomaterials on cardiac cell adhesion, protein expression (i.e. maturation), and tissue-level excitability. Our results demonstrated that either embedding nanomaterials (i.e. GNRs and SNPs) or increasing the matrix stiffness significantly promoted cellular retention and the expression of cardiac-specific markers, including sarcomeric α-actinin (SAC), cardiac troponin I (cTnI) and connexin43 (Cx43) gap junctions. Notably, excitation voltage thresholds at a high frequency (i.e. 2 Hz and higher), in both coupled and uncoupled gap junctions induced by heptanol, were lower for scaffolds embedded conductive GNRs or non-conductive SNPs, independent of matrix stiffness. Overall, our findings demonstrated that the sole presence of nanomaterials within the scaffolding matrix had a more pronounced influence as compared to the scaffold stiffness on the cell–cell coupling, maturation and excitability of engineered cardiac tissues.


Introduction

During the past decade, electrically conductive nanocomposite scaffolds have been at the center of attention for engineering functional cardiac tissues with a mature phenotype, enhanced electrical excitability and signal propagation.1–8 Conductive nanocomposite scaffolds are typically fabricated by incorporating electrically conductive nanomaterials, such as gold nanomaterials (GNMs),2,3 carbon nanotubes (CNTs),5,9 and reduced graphene oxide (rGO),6 within the macroporous matrix of scaffolding biomaterials, including hydrogels or electrospun nanofibers.10–12 The addition of conductive nanomaterials has been argued to facilitate electrical signal propagation within the scaffold via bridging the insulating matrix pore walls, resulting in a biomimetic matrix similar to the native heart extracellular matrix (ECM), leading to enhanced electrical coupling and excitability of cardiomyocytes (CMs).2,3,5,9 In addition, as a secondary influence, conductive nanocomposite scaffolds have been shown to promote cellular adhesion and retention2,13,14 as well as the expression of cardiac-specific proteins.1,3,5,15–19

GNMs and CNTs have been among the most commonly used nanomaterials for fabricating conductive cardiac tissues.1–3,5,20–22 For instance, in a pioneering study by Dvir,3 conductive cardiac tissues were fabricated by seeding CMs on GNM-embedded alginate hydrogels. Alginate–GNM hydrogels enhanced the maturation of CMs, manifested by upregulated expressions of sarcomeric α-actinin (SAC), cardiac troponin I (cTnI), and connexin43 (Cx43) gap junctions. In addition, external electrical stimuli were propagated globally (i.e. synchronicity) only within the alginate–GNM hydrogel, while pure alginate did not facilitate a similar signal transmittance. In another work by Shin,5 gelatin–CNT hydrogels were synthesized to engineer conductive cardiac tissues. Incorporation of CNTs within photocrosslinkable gelatin methacrylate (GelMA) hydrogels significantly increased SAC, cTnI and Cx43 expressions and induced the formation of striated sarcomeres. Furthermore, GelMA–CNT hydrogels enhanced the excitability (i.e. excitation voltage threshold) and synchronized the spontaneous contractility of CMs. In these studies, it was rationalized that CMs incorporated within macroporous matrix of pristine scaffold (i.e. without the presence of conductive nanomaterials) are electrically insulated by non-conductive polymeric pore walls, which can result in discrete cell–cell electrical coupling, reduced conduction velocity and unsynchronized contractions throughout the engineered tissue.

Aside from electrical conductivity, the incorporation of nanomaterials has been also shown to influence mechanical stiffness23,24 and generate nano-scale surface topographies within the scaffolding biomaterials.25,26 With respect to cardiac tissue engineering, enhancing the mechanical stiffness and topographical characteristics of the substrate, without the use of nanomaterials, has been shown to promote maturation as well as contractile and electrical functionalities of both primary and stem cell-derived CMs.27–32 For example, CMs seeded on polyacrylamide hydrogels with higher stiffness (due to the increased prepolymer concentration) demonstrated enhanced expressions of cardiac proteins and contractile forces in comparison with CMs cultured on softer polyacrylamide hydrogels.27 However, despite these reports, what is missing in the literature is whether the sole presence of nanomaterials, regardless of electrical conductivity, could lead to synchronized electrical and contractile functionalities among CMs.

In our recent studies,1,2 we developed gold nanorod (GNR)-embedded GelMA nanocomposite hydrogels for engineering conductive cardiac tissues. GNRs significantly increased the conductivity and stiffness of the hydrogel matrix. Additionally, scanning electron microscopy (SEM) micrographs demonstrated the localization of GNRs on the surface of the GelMA matrix, inducing nano-scale topographies.3,33 GelMA–GNR hydrogels significantly increased cellular retention, maturation (e.g. protein expression and sarcomere formation), and electrical excitability within the engineered tissues. In this study, we aimed to take a significant step forward and developed four different classes of gelatin-based scaffolds with embedded conductive GNRs as well as non-conductive silica nanoparticles (SNPs) to dissect the role of matrix stiffness and the presence of nanomaterials, regardless of their electrical conductivity, on the functionalities of cardiac tissues. Specifically, through the synthesis of a GelMA–SNP nanocomposite hydrogel, featuring a non-conductive and mechanically soft matrix, we were able to investigate that to what extent the sole presence of nanoparticles influences the maturation and excitability of the engineered cardiac tissues. To our knowledge our study is the first to compare side-by-side the influence of electrically conductive and non-conductive nanocomposite scaffolds (i.e. with embedded conductive and non-conductive nanoparticles) on the maturation and excitability of engineered cardiac tissues.

Experimental methods

Materials

Gold(III) chloride trihydrate (HAuCl4), sodium borohydride (NaBH4), silver nitrate (AgNO3), hexadecyltrimethylammonium bromide (CTAB), ascorbic acid, gelatin (Type A), methacrylic anhydride, 3-(trimethoxysilyl) propyl methacrylate (TMSPMA), and 2-hydroxy-1-(4-(hydroxyethoxy) phenyl)-2-methyl-1-propanone (photoinitiator) were purchased from Sigma Aldrich and used without any modifications. 80 nm silica nanoparticles (SNPs) with the specific surface area of 30–40 m2 g−1 and a skeletal density of 2.1–2.2 g cm−3 were purchased from General Engineering and Research (San Diego, USA).

Synthesis of GelMA hydrogels and GNRs

GelMA and GNRs were synthesized based on the previously established protocols.2,34,35 To synthesize GelMA, 10% wt/v type A gelatin was dissolved in phosphate buffer saline (PBS) and the resulting mixture was stirred at 50 °C for 1 h. Afterwards, the gelatin solution was methacrylated (high degree) by drop-wise addition of methacrylic anhydride (8% wt/v). After 3 h, the resulting solution was diluted 5 times by introducing 50 °C PBS to stop the methacrylation reaction. The final solution was then dialyzed (12–14 kDa cutoff) against deionized water at 50 °C for 7 days to eliminate unreacted compounds and salts. Lastly, the purified gelatin solution was filtered, lyophilized and stored at −80 °C. GNRs with an aspect ratio of approximately 4 (length 60 nm and width 15 nm) were synthesized through anisotropic growth of gold nanoparticles (GNPs). First, GNPs were synthesized by adding 240 μL ice-cold NaBH4 (10 mM) to an aqueous solution of HAuCl4 (2 mL, 0.5 mM) and CTAB (2 mL, 0.2 M), and vortexed for 1 min. Afterwards, 48 μL of GNPs (seed solution) was introduced into the growth medium, containing HAuCl4 (20 mL, 1 mM), CTAB (20 mL, 0.2 M), AgNO3 (1.12 mL, 4 mM) and ascorbic acid (280 μL, 78.8 mM), and the whole mixture was kept undisturbed at 30 °C overnight. The synthesized GNRs were purified by centrifugation (12[thin space (1/6-em)]000 rpm for 10 min) and dispersed in deionized water. The formation of GNRs was confirmed using transmission electron microscopy (TEM) operating at 200 kV accelerating voltage.

Fabrication and characterization of pristine GelMA, SNP- and GNR-embedded GelMA nanocomposite hydrogel constructs

Hydrogel constructs were fabricated based on UV photocrosslinking.2,34 Briefly, a clear solution of GelMA hydrogels (5% and 20% wt/v) was prepared in 0.5% (wt/v) photoinitiator in PBS. Next, GNRs (1.5 mg mL−1) and SNPs (9.69 mg mL−1) were mixed with the GelMA prepolymer solution (5% wt/v) and sonicated (bath) for 1 h to create homogeneous colloid mixtures. SNPs were ultrasonicated (on ice) in deionized water for at least 2 h prior to mixing with the GelMA prepolymer solution. We chose 80 nm (diameter) SNPs for our study as they provided approximately similar morphology and size compared to the synthesized GNRs with ∼60 nm length and ∼20 nm width. The concentration of SNPs was determined to provide an approximately similar number of nanoparticles (1.36 × 1013 particles per mL) in comparison with a GNR concentration of 1.5 mg mL−1. Specifically, 1.5 mg mL−1 of our synthesized GNRs contains 1.36 × 1013 particles per mL. In addition, silica has a density of 2.65 g cm−3 and the volume of one SNP with a diameter of 80 nm is 2.68 × 10−16 cm3. Therefore, 9.69 mg mL−1 of SNPs is required to provide 1.36 × 1013 particles per mL concentration similar to the concentration of GNRs.

The UV exposure time for each hydrogel group was optimized in order to fabricate fully crosslinked and formed hydrogel constructs. There are parameters influencing the UV exposure time, for instance, the hydrogel prepolymer solution, construct thickness and hydrogel composition (e.g. with embedded nanomaterials and concentration of gel).36 For example, increasing the concentration of the prepolymer solution (e.g. from 5% to 20% wt/v) has been shown to increase the polymer chain density, eventually leading to higher UV exposure time in order to fully crosslink the hydrogel constructs.37 In addition, hydrogel constructs with a higher thickness require a higher exposure time as compared to the thinner ones in order to allow full crosslinking of the whole construct. Finally, embedding nanomaterials (e.g. GNRs and SNPs), which absorb UV light, within the hydrogel matrix might delay the initiation of the crosslinking process and consequently lead to a higher exposure time.26,38,39 Therefore, based on these factors and assessments in our study, the UV exposure times for the 150 μm thick hydrogel construct were optimized to be: GelMA (5%) = 6 s, GelMA (20%) = 10 s, GelMA–SNP = 10 s and GelMA–GNR = 30 s.

After UV crosslinking, the fabricated hydrogel constructs were fully hydrated for at least 2 h in PBS before performing material characterization. The electrical conductivity and mechanical stiffness of the fabricated hydrogel constructs were measured using atomic force microscopy (AFM) and an LCR meter, based on the established protocols described in our previous studies.1,2,35 Specifically, hydrated hydrogel constructs (n = 6) were placed between two glass slides with a conductive indium tin oxide coating (Sigma Aldrich) and connected to an LCR meter device (Agilent 4284A) and the impedance was measured at different AC bias frequencies (10 Hz to 1 MHz). The mechanical stiffness of the hydrated hydrogel constructs (n = 3) was evaluated by measuring Young's modulus using AFM (MFP-3D AFM, Asylum Research) with silicon nitride probes (MSNL, Bruker) with a nominal spring constant of 0.1 N m−1 and an effective half cone angle of 15° at 37 C°. The Young's moduli were obtained by fitting force–indentation curves to the Sneddon contact model for a conical indenter.40

CM isolation and culture of cardiac tissue constructs

All hydrogel constructs were sterilized by washing two times (10 min intervals) in an antibiotic solution (2% penicillin/streptomycin). The samples were then washed twice using culture media prior to CM seeding. CMs were isolated from the ventricular region of 2-day old neonatal rats based on the previously established protocols.34 Briefly, hearts were excised from the rats and cut into small pieces (1–2 mm3). The sliced tissues were then treated with trypsin without EDTA (0.05% v/v in Hank's balanced salt solution) overnight at 4 °C. Afterwards, the tissues were subjected to serial (3–5 times) collagen digestion (type 2, 1 mg mL−1). The isolated cardiac cells were filtered using a cell strainer (70 μm) and incubated for 1 h at 37 °C to allow the separation of CMs and Cardiac Fibroblasts (CFs). Cardiac tissue constructs were fabricated by seeding CMs on the hydrogel constructs (7.5 × 105 per construct) for 1 day and culturing for 7 days in an incubator at 37 °C with 5% CO2. The culture media, containing Dulbecco's modified Eagle's medium supplemented with 10% fetal bovine serum, 1% glutamine, and 1% penicillin/streptomycin, were exchanged every other day.

Characterization of CM retention and expression of cardiac markers

Adhesion and retention (n = 6) of the seeded CMs were evaluated by measuring the area percentage of hydrogel constructs covered by tightly attached cells on day 1. Specifically, the samples were washed two times with pre-warmed culture media to eliminate loosely adhered cells prior to phase-contrast imaging. At least 5 technical replicates were examined for the cell retention measurement. The phenotypic maturation of engineered tissues was assessed by immunofluorescent staining of cardiac-specific markers, including SAC, cTnI and Cx43 on day 7. Briefly, the tissue constructs were first fixed using 4% paraformaldehyde for 20 min, permeabilized with 0.1% Triton X-100 for 10 min, and finally blocked by 10% goat serum for 1 h at room temperature. Afterwards, the tissue constructs were incubated with primary antibodies for SAC (Sigma Aldrich, USA), cTnI and Cx43 (Abcam, USA) overnight at 4 °C, followed by fluorescent staining using secondary antibodies for 1–2 h at room temperature (Life Technologies, USA). Z-Stacked immunostained images were captured using a fluorescence microscope (Zeiss ObzerverZ1, USA) equipped with ApoTome2 and were analyzed using NIH ImageJ for measuring the protein expression (n > 3). The area of the fluorescence signal (μm2) was calculated for each protein of interest and normalized by dividing with the total number of nuclei.

External electric field simulation and tissue excitability assessment

A custom-made field stimulation chamber was used to apply electrical stimuli to the hydrogel tissue constructs as previously described in the work of ours and others.2,41,42 Briefly, two carbon rod electrodes with 1 cm spacing were glued inside a 60 mm plastic dish using an electrically insulating silicone adhesive. The carbon electrodes were connected to a function generator (BK PRECISION) using platinum wires. Electrical field pulses with 2–5 ms duration at different frequencies (0.5, 1, 2, 3, 4, and 5 Hz) were used for the experiment. The engineered tissues on day 7 of culture were evaluated for the excitation voltage threshold, which was defined as the minimum voltage potential required to induce global synchronous contractions. The samples were placed in the stimulation chamber and the voltage potential was increased step-wise by 0.2 V until tissue-level synchronous contractions of CMs were evident. For the heptanol treatment experiment, tissue constructs were incubated (37 °C) with 2 mM heptanol, as a gap junction uncoupler agent,5,43,44 for 20 min and the excitation voltage threshold was measured at 0.5, 1, 2 and 3 Hz frequencies. All the measurements were performed at 37 °C and the pH was maintained at approximately 7 by adding 10 mM HEPES buffer to the cell culture media. Video microscopy was carried out using an inverted Zeiss microscope (ObzerverZ1, USA) and the captured movies were analyzed using a custom-written MATLAB code45 to extract beating signals.

Statistical analysis

All the collected data were plotted using GraphPad Prism. The statistical analysis was performed based on Student's t-test and one-way and two-way ANOVA followed by a post-hoc Tukey multiple comparison test. The results were demonstrated as mean ± standard deviation (SD). The p value less than 0.05 (two-sided) was considered as the statistical significance across the experimental groups.

Results and discussion

Development of hydrogel constructs with specific matrix characteristics

We developed four hydrogel candidates, including GelMA (5%), GelMA (20%), GelMA–SNP, and GelMA–GNR, with controlled matrix properties to conduct our studies. The pristine GelMA (5%) hydrogel, without the presence of nanomaterials, was selected as the negative control group with low mechanical stiffness (i.e. soft matrix) and low electrical conductivity, while the GelMA (5%) hydrogel embedded with 1.5 mg mL−1 GNRs (GelMA–GNR) was used as the hydrogel candidate (positive control) featuring high mechanical stiffness, high electrical conductivity and nano-scale surface topographies, consistent with our previous work.1,2 The outcomes of our previous studies showed that the GelMA hydrogels incorporating 1.5 mg mL−1 of GNRs demonstrated excellent cell adhesion and retention affinity, cardiac maturation, excitability and contractility. Therefore, the selection of 1.5 mg mL−1 of GNR candidate makes it an appropriate positive control group to conduct our study. In addition, we embedded SNPs (Fig. 1A) within GelMA (5%) to develop a hydrogel candidate with insulating (i.e. poor electrical conductivity46,47) and soft matrix properties, however with the presence of nano-scale topographies.48–50 We chose 80 nm (diameter) SNPs as they exhibited similar morphology and size compared to the synthesized GNRs (∼60 nm length and ∼20 nm width) (Fig. 1B). Furthermore, the concentration of SNPs was set to 9.69 mg mL−1 to provide a similar number of nanoparticles (1.36 × 1013 particles per mL) as compared to the GelMA–GNR hydrogel candidate (1.5 mg mL−1). Lastly, the pristine GelMA hydrogel with a 20% (v/v) prepolymer concentration was introduced as the hydrogel group with a similar mechanical stiffness (i.e. high stiffness) as compared to GelMA–GNR (1.5 mg mL−1), however with low electrical conductivity and without the presence of nanomaterials (i.e. lack of nano-scale matrix topographies). With this selection of four hydrogel candidates (i.e. GelMA (5%), GelMA (20%), GelMA–SNP and GelMA–GNR) we were able to modulate three different matrix properties, namely: mechanical stiffness, electrical conductivity and nano-scale matrix topography.
image file: c8bm01050a-f1.tif
Fig. 1 (A) SEM images of SNPs51 and (B) TEM images of GNRs showing the morphology of nanomaterials. (C) Impedance measurements demonstrating the enhanced electrical conductivity of GelMA–GNR and high electrical resistance of GelMA–SNP hydrogel constructs. (D) Young's moduli of all fabricated hydrogel constructs showing the increased mechanical stiffness of GelMA–GNR and GelMA (20%) in contrast to the soft matrices of GelMA (5%) and GelMA–SNP hydrogels (*p-value < 0.05).

The fabricated hydrogel constructs, including GelMA (5%), GelMA (20%), GelMA–SNP and GelMA–GNR, were characterized for mechanical stiffness and electrical conductivity. The impedance measurement demonstrated (Fig. 1C) that the inclusion of GNRs significantly reduced the electrical resistance of GelMA–GNR (2.5 ± 0.03 kΩ at 20 Hz) as compared to both pristine GelMA (5%) (12.65 ± 5.21 kΩ at 20 Hz) and GelMA (20%) (21.58 ± 3.56 kΩ at 20 Hz) hydrogels. Similar findings were also reported in previous studies of ours1,2 and others.3–5,15,16 The enhanced conductivity is attributed to the induced resistive current through GNRs within the insulating matrix of the polymeric scaffold. On the other hand, GelMA–SNP hydrogels demonstrated the highest impedance (30.77 ± 12.36 kΩ at 20 Hz) compared to pristine and GNR-embedded GelMA hydrogels, indicating the insulating property of the GelMA–SNP scaffold due to the presence of SNPs.47 With respect to the mechanical stiffness of the hydrogel candidates, Fig. 1D showed similar Young's moduli for GelMA (20%) (1.63 ± 0.24 kPa) and GelMA–GNR (1.37 ± 0.43 kPa). Furthermore, the addition of 9.69 mg mL−1 SNPs did not lead to a significant change in the mechanical stiffness of GelMA–SNP (0.52 ± 0.09 kPa for GelMA (5%) and 0.55 ± 0.17 kPa for GelMA–SNP). These results are consistent with previously reported studies for polyethylene glycol (PEG) hydrogels incorporated with 1% (wt/v) SNPs.26

Impact of hydrogel matrix characteristics on CM retention and protein expression

Fig. 2 displays the phase-contrast images and measured retention of CMs on the different hydrogel constructs on day 1. The seeded CMs on the hydrogel constructs demonstrated sparse and aggregated adhesion patterns on GelMA (5%), while more compact and uniform adhesion patterns were observed on GelMA (20%), GelMA–SNP and GelMA–GNR hydrogels. The quantified data, presented in Fig. 2B, demonstrated that CMs covered approximately 30.9% of GelMA (20%), 68.8% of GelMA–SNP and 73.6% of GelMA–GNR hydrogel constructs, which were all significantly higher than the 14.2% cell retention on the GelMA (5%) construct. Overall, for hydrogels not modified with nanomaterials (GelMA (20%)), an increasing trend of cellular retention was seen by enhancing the hydrogel stiffness. However, incorporating non-conductive nanomaterials (SNPs) within the matrix showed a significantly higher cellular retention despite a similar Young's modulus as compared to the soft GelMA (5%). Lastly, electrically conductive GelMA–GNR hydrogels also substantially increased the CM adhesion affinity similar to the non-conductive GelMA–SNP.
image file: c8bm01050a-f2.tif
Fig. 2 (A) Phase-contrast images of adhered CMs on GelMA (5%), GelMA (20%), GelMA–SNP and GelMA–GNR hydrogel constructs on day 1 showing (B) a significant increase in CM adhesion and retention due to the enhanced scaffold stiffness and the incorporation of SNPs and GNRs (*p-value < 0.05).

In our previous work, we showed that GNRs are located on the pores of the GelMA hydrogel matrix,2 which has been shown to induce nano-scale surface topographies.33 Therefore, these findings primarily support the superior impact of the matrix nano-scale topography, due to the embedded nanomaterials, on cell adhesion and retention as compared to the matrix stiffness. In addition, previous studies have also demonstrated that hydrogel matrices either with high stiffness or incorporated with nanomaterials26,48 promoted cellular adhesion and retention.27,30,52 In this regard, the enhanced cell adhesion affinity is attributed to the increased expressions of integrin and mechanosensitive kinases, such as focal adhesion kinase (FAK), extracellular signal-regulated kinases (ERK), and protein kinase B (AKT).53 Furthermore, nano-scale topographies formed by the embedded nanomaterials could also increase the surface area and protein adsorption and generate cell anchoring sites, leading to improved cell adhesion and retention.48,54 The CM retention on the conductive GelMA–GNR hydrogel was similar to that of the non-conductive GelMA–SNP hydrogel, suggesting the negligible impact of matrix electrical conductivity on cellular adhesion and retention.

Next, we assessed the expression of cardiac-specific markers to determine the impact of scaffold matrix properties (electrical, mechanical and topographical) on the phenotypic maturation of engineered cardiac tissues. Fig. 3A shows the immunofluorescence images of cardiac-specific markers, including SAC, cTnI and Cx43 gap junctions. A clear formation of striated sarcomere structures (SAC staining, white arrows) was evident on GelMA (20%), GelMA–SNP and GelMA–GNR tissue constructs as compared to the disoriented SAC expressions on the pristine GelMA (5%) hydrogel. In addition, SAC stained images exhibited more elongated and uniaxially aligned sarcomere structures on GelMA (20%) and SNP- and GNR-embedded tissue constructs in contrast to the random and round morphology (white arrows) on GelMA (5%). The expression of cTnI also followed a similar pattern as SAC. GelMA (20%), GelMA–SNP and GelMA–GNR cardiac tissues illustrated extended and abundant expressions of cTnI, while scattered patterns were observed on GelMA (5%) hydrogel tissues. Cx43 gap junctions (red punctuated stains) were present among CMs on all hydrogel constructs. However, a less homogeneous distribution was evident on GelMA (5%) as compared to the other hydrogel groups. Similar results have also been reported in other studies demonstrating the mature phenotype of primary and iPSC-derived CMs.17,55 In addition to the representative images, the quantified fluorescent area (Fig. 3B–D) of cardiac specific proteins also confirmed significantly higher expressions of SAC, cTnI and Cx43 on high stiffness GelMA (20%) as well as nanomaterial-embedded GelMA–SNP and GelMA–GNR matrices in comparison with GelMA (5%). Moreover, we did not find substantial differences in terms of protein expression between electrically conductive GelMA–GNR and non-conductive GelMA–SNP hydrogels. Overall, our findings suggested that enhancing the stiffness or embedding nanomaterials within scaffolding matrices promoted the maturation of engineered cardiac tissue, regardless of the electrical conductivity of the matrix.28,31,56,57


image file: c8bm01050a-f3.tif
Fig. 3 (A) The Z-stacked immunostained images and (B) the positive fluorescence area of SAC, Cx43 and cTnI expressions on day 7, illustrating enhanced cardiac maturation via improving the scaffold mechanical, topographical and electrical properties. White arrow heads indicate the round and disconnected morphology of sarcomere structures on GelMA (5%) in contrast to the intact and uniaxially aligned sarcomeres (i.e. higher maturation) on GelMA (20%), GelMA–SNP and GelMA–GNR hydrogels. All scale bars represent 20 μm (*p-value < 0.05).

Influence of hydrogel matrix characteristics on the electrical excitability of engineered cardiac tissues

External electric field stimulation has been utilized as a method to evaluate the tissue-level excitability and contractility of engineered cardiac tissues.4,5,41 In our work, we fabricated and utilized a custom-made stimulation chamber1,2,22,42 to induce contractions and simultaneously measure the excitability of the cardiac tissues. Fig. 4 exhibits the excitation voltage thresholds and representative extracted beating signals for the four groups of hydrogel candidates, GelMA (5%), GelMA (20%), GelMA–SNP and GelMA–GNR tissues, at different stimulation frequencies.
image file: c8bm01050a-f4.tif
Fig. 4 (A–D) The extracted CM beating signals in response to electrical stimuli at 0.5–4 Hz, showing a higher beating capture rate for GelMA–SNP and GelMA–GNR (4 Hz) as compared to GelMA (5%) (2 Hz) and GelMA (20%) (3 Hz). (E) Excitation thresholds of pristine GelMA and embedded tissue constructs following electrical field stimulation (0.5–4 Hz), illustrating the enhanced excitability of cardiac tissues due to the incorporation of SNPs and GNRs (*p-value < 0.05).

As can be seen, all the hydrogel groups were able to fully accommodate electrical stimulation frequencies up to 2 Hz. In this frequency range, GelMA (5%) cardiac tissues showed the highest excitation voltage threshold (∼8.4 V) as compared to GelMA (20%) (∼4.3 V), GelMA–SNP (∼3.8 V) and GelMA–GNR (∼3.7 V). In addition, our results did not show any significant differences for excitation voltage thresholds across the GelMA (20%), GelMA–SNP and GelMA–GNR groups. By increasing the frequency to 3 and 4 Hz, however both soft GelMA (5%) and stiff GelMA (20%) hydrogel tissues were no longer able to follow the stimulation regime and tissue contractions became chaotic (Fig. 4A, B and ESI Movies M1 and M2). On the other hand, GelMA–SNP and GelMA–GNR tissue constructs demonstrated contractile behavior consistent with the induced stimulation frequencies at 3 and 4 Hz (Fig. 4C, D and ESI Movies M3 and M4). None of the hydrogel groups could fully accommodate electrical stimulations with frequencies larger than 4 Hz (data not shown).

To further elucidate the potential compensatory impact of matrix electrical conductivity on the excitability of engineered cardiac tissues, we utilized heptanol to diminish the intrinsic electrical coupling within CMs through Cx43 gap junctions. Heptanol has been widely used to uncouple electrical gap junctions, such as Cx43, which leads to reduced conduction velocity and unsynchronized contractions in CMs.44,58,59 Excitation voltage thresholds of the heptanol-treated cardiac tissue are illustrated in Fig. 5 for 0.5, 1 and 2 Hz frequencies. As can be seen, after 20 min of exposure to 2 mM heptanol, GelMA (5%) tissue constructs were no longer able to follow the electrical stimulation at any of the applied frequencies (Fig. 5A and E and ESI Movie M5). In addition, CMs cultured on GelMA (20%) hydrogels only responded to stimuli with 0.5 and 1 Hz upon heptanol exposure and were not able to accommodate 2 Hz stimuli (Fig. 5B and E and ESI Movie M6). On the other hand, both GelMA–SNP and GelMA–GNR cardiac tissues fully accommodated up to 2 Hz electrical stimuli even in the presence of heptanol (Fig. 5C–E and ESI Movies M7 and M8). Lastly, none of the engineered cardiac tissues were able to generate contractions upon electrical stimulation with frequencies higher than 2 Hz (data not shown). When a higher concentration of heptanol (3 mM) was tested, we were not able to induce contractions on any of the cardiac tissue contracts at any frequencies with a voltage potential up to 10 V.


image file: c8bm01050a-f5.tif
Fig. 5 (A–D) The extracted beating signals of stimulated cardiac tissues upon 20 min of treatment with heptanol (2 mM), demonstrating that GelMA–SNP and GelMA–GNR hydrogels accommodated electrical stimuli with higher frequencies (up to 2 Hz) in comparison with GelMA (20%) (up to 1 Hz). (E) Excitation voltage thresholds of engineered tissues after 20 min of heptanol (2 mM) exposure, showing higher excitability of SNP- and GNR-embedded constructs as compared with 5% and 20% GelMA constructs.

One of the main focuses of our work was to provide more insights into the presence of nanomaterials, matrix stiffness and electrical conductivity on the excitability and contractility of the engineered cardiac tissues. In this regard, incorporating conductive GNRs within the hydrogel matrix enhanced mechanical stiffness and generated nano-scale topographies. Along with GelMA–GNR, we also fabricated two other non-conductive hydrogel candidates to conduct our study: a GelMA (20%) hydrogel with high matrix stiffness but without the presence of nanomaterials (i.e. lack of nano-scale topographies), and a GelMA–SNP hydrogel with insulating property, low stiffness but with nano-scale topographies due to the presence of SNPs. Our results of electrical stimulation (Fig. 4 and 5), in both settings of coupled and uncoupled gap junctions using heptanol, suggested that the sole presence of nanomaterials within the scaffolding matrix, either conductive (i.e. GNRs) or non-conductive (i.e. SNPs), and the induced nano-scale topographies have a more pronounced effect as compared to the matrix stiffness on the excitability and contractility of engineered cardiac tissues. In other words, when embedding nanomaterials in hydrogel matrices, scaffold conductivity may not be the sole player influencing the excitability of the engineered cardiac tissues. As has been also shown in previous studies of ours and others2,48,54 and discussed earlier, nano-scale topographies formed by embedded nanomaterials can increase the surface area and protein adsorption and generate cell anchoring sites within the hydrogel matrix, leading to improved cell adhesion, retention and expression of integrins. On the other hand, the enhanced cell–matrix interactions and expression of integrins have been shown to significantly improve protein expression and maturation in CMs.27–32 Our studies presented herein are indeed consistent with previous work, where we showed that the enhanced cellular retention (Fig. 2) was accompanied by enhanced cardiac cell–cell junctions and protein expression on the nanocomposite hydrogels (Fig. 3). Therefore, we rationalize that the sole presence of nanomaterials, either conductive GNRs or non-conductive SNPs, within the GelMA matrix and the resulting nano-scale topographies, leads to enhanced cellular retention which further promotes cell–cell coupling through gap junctions as well as protein expressions. The enhanced cell–cell coupling and protein expressions in turn result in pronounced maturation and contractility (i.e. lower excitation threshold and higher maximum capture rate) of the engineered tissues. We also speculate that the presence of nanomaterials within the matrix of the scaffold can result in other unique interactions with cells, for example through the adsorbed proteins on the surface of nanomaterials,60–62 that have been shown to exert a significant impact on the functionalities of CMs.63–65

Despite the significance of our study, there are other techniques that can be utilized to further improve our mechanistic understanding of cell–matrix interactions and the role of conductivity in the functionalities of engineered cardiac tissues. For example, in our study we focused on the tissue-level electrical excitability of engineered cardiac constructs via utilizing an electric field stimulation chamber. However, there are still other critical factors that may influence the excitability and electrical coupling of CMs, such as membrane depolarization threshold and conduction velocity,66,67 which need to be investigated in the future. In addition, the utilization of other characterization techniques, such as voltage sensitive fluorescent dyes and point electrical stimulation using bipolar microelectrodes, may provide more insight on the differences between conductive and non-conductive nanomaterials on the electrophysiological properties (e.g. conduction velocity) of engineered cardiac tissues. Lastly, we utilized neonatal ventricular CMs isolated from 2-day old rats for our study, which are able to transition from the neonatal to mature electrophysiological state, featuring robust expression of Cx43 electrical gap junctions, stable membrane resting potential, and enhanced depolarization amplitude.68–70 Such intrinsic (i.e. native) cellular properties may shield the potential impact of scaffold conductivity on the excitability and synchronicity of the engineered tissue. In this regard, utilization of other cell types, such as induced pluripotent- and embryonic stem cell-derived CMs, which exhibit less electrophysiological maturation,71–73 may enable us to better dissect the role of scaffold conductivity in the excitability and synchronicity of engineered cardiac tissues.

Conclusions

Our goal in this study was to investigate the role of mechanical stiffness and nano-scale topography of the scaffolding matrix on cell–cell coupling, maturation and electrical excitability of engineered cardiac tissues, with different electrical conductivity. We designed and fabricated four different hydrogel candidates, including GelMA (5%) (mechanically soft), GelMA (20%) (mechanically stiff), GelMA–SNP (non-conductive with nano-topographies and mechanically soft) and GelMA–GNR (conductive with nano-topographies and mechanically stiff). Our results demonstrated that GelMA–SNP and GelMA–GNR hydrogels significantly improved CM adhesion affinity as compared to pristine GelMA (5% and 20%), highlighting the influence of nano-scale topography, due to the presence of nanomaterials, on cellular adhesion and retention, independent of the elastic modulus of scaffolds. Furthermore, protein expression results revealed that the increase in hydrogel stiffness alone promoted maturation of engineered cardiac tissues; however, such enhancement was also evident in the presence of SNP and GNR nanomaterials. The induced stiffness and nano-scale topography significantly enhanced electrical excitability of engineered cardiac tissues as shown by the decreased excitation voltage thresholds. Specifically, GelMA–SNP and GelMA–GNR cardiac tissues accommodated external electrical stimuli at higher frequencies (i.e. higher capture rate) in both settings of coupled and uncoupled gap junctions. Most importantly, similar cellular retention, cell–cell coupling (i.e. gap junctions), cardiac-specific protein expression and excitation threshold were observed between conductive GelMA–GNR and non-conductive GelMA–SNP tissue constructs. These findings suggest the prominent role of the nano-scale surface topography of nanocomposite scaffolds, due to the presence of nanomaterials within the hydrogel matrix, compared to mechanical stiffness in cardiac cell–cell coupling, maturation and functionalities of the engineered tissue.

Conflicts of interest

The authors declare no conflict of interest.

Acknowledgements

We would like to acknowledge the generous support from Phoenix Children's Hospital (PCH) Leadership Circle Award. We would like to acknowledge Shiyi Liu (Dr Chae Lab) for his help on impedance measurement.

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Footnote

Electronic supplementary information (ESI) available. See DOI: 10.1039/c8bm01050a

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