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. Author manuscript; available in PMC: 2009 Aug 27.
Published in final edited form as: Prog Biophys Mol Biol. 2008 Feb 16;97(2-3):367–382. doi: 10.1016/j.pbiomolbio.2008.02.017

Cell cultures as models of cardiac mechanoelectric feedback

Yibing Zhang 1, Rajesh B Sekar 1, Andrew D McCulloch 2, Leslie Tung 1
PMCID: PMC2733372  NIHMSID: NIHMS59773  PMID: 18384846

Abstract

Although stretch-activated currents have been extensively studied in isolated cells and intact hearts in the context of mechanoelectric feedback (MEF) in the heart, quantitative data regarding other mechanical parameters such as pressure, shear, bending, etc, are still lacking at the multicellular level. Cultured cardiac cell monolayers have been used increasingly in the past decade as an in vitro model for the studies of fundamental mechanisms that underlie normal and pathological electrophysiology at the tissue level. Optical mapping makes possible multisite recording and analysis of action potentials and wavefront propagation, suitable for monitoring the electrophysiological activity of the cardiac cell monolayer under a wide variety of controlled mechanical conditions. In this paper, we review methodologies that have been developed or could be used to mechanically perturb cell monolayers, and present some new results on the acute effects of pressure, shear stress and anisotropic strain on cultured neonatal rat ventricular myocyte (NRVM) monolayers.

Keywords: Mechanoelectrical feedback, cardiac cell monolayer, optical mapping, electrophysiology, shear stress, pressure, anisotropic stretch

1. Introduction

Composed of electrically excitable cells, the heart works as a mechanical pump of blood for the lifetime of the human body. In the closed mechanoelectrochemical transduction loops inside the cardiac cell, mechanoelectrical feedback (MEF) — also referred to as mechanoelectric coupling, contraction-excitation coupling, reverse excitation-contraction coupling or mechanically mediated crosstalk (Lab, 2005) — plays an important role in its physiological and pathological functions. From the classical viewpoint of a feedback loop, a mechanical disturbance in the environment produces a change in length and tension of the cardiac cell, which feeds back and alters the excitation of the cell that controls mechanical contraction and its associated length and tension (Kohl et al., 1999).

Studies of MEF traditionally focus on responses of cardiac cells to membrane stretch or tissue to wall stretch (Kohl and Ravens, 2003). However, mechanical loading of cardiac tissue in vivo is three-dimensional and multi-scale, depending on heterogeneous wall stress distributions (Mihailescu and Abel, 1994) that give rise to complex wall deformations such as transverse shearing between the laminar sheets of myocardial cells (Costa et al., 1999). These three-dimensional deformations also give rise to fluid shear stresses that result from the flow of interstitial fluids around cells (Lorenzen-Schmidt et al., 2006). In single ventricular myocytes, axial stretch, local indentation, and hypoosmotic swelling affect membrane currents in different ways (Isenberg et al., 2003; Sasaki et al., 1992), One of the impediments of the study of MEF at the cellular level is the technical complexity of the experiments (Kohl and Ravens, 2003), which involve application of a mechanical perturbation (the input) while monitoring a change in electrophysiology (the readout).

Simplified cell culture models can circumvent complexities of conventional tissue models and provide an in vitro preparation that can be tested under controlled conditions. Electrophysiological readouts of cellular voltage can be obtained by microelectodes (Jongsma and van Rijn, 1972), multielectrode arrays (Israel et al., 1990), or optical mapping (Fast and Kleber, 1993; Rohr and Salzberg, 1994). Optical mapping makes possible the simultaneous measurement of action potentials from a large number of recording sites, which can be used to generate maps of electrical propagation. Readouts of cellular calcium can also be obtained by optical mapping (Fast, 2005). Excellent reviews of optical mapping approaches on cardiac tissue have been published for intact tissue (Efimov et al., 2004) and cultured cell monolayers (Entcheva and Bien, 2006). Compared with microelectrode array measurements of cardiac electrical activity, optical mapping not only provides higher spatial resolution without recording artifacts which can arise from electrical stimulators, but can readily be integrated into a test system where different mechanical perturbations are applied to the cardiac tissue. Quantitative information relevant to the mechanisms of MEF under different mechanical stimuli, such as pressure, shear stress and anisotropic strain can thus be obtained.

This paper reviews methodologies that have been developed to mechanically perturb cardiac cell monolayers (section 2), and presents effects of acute changes in hydrostatic pressure, shear stress and anisotropic stretch on cardiac cell monolayers (section 3), followed by discussion (section 4) and conclusions (section 5).

2. Mechanical stimulation of cardiac monolayers

2.1 Why use a cultured cell monolayer?

Presently, experimental models for the study of MEF are generally either the intact tissue (including whole heart) or single cells. In situ, cells are mechanically coupled to the extracellular matrix via integrins and to other cells via cell-cell adhesion molecules (Ingber, 1997). Given that cells in situ experience anisotropic stretch, shear strains and heterogeneous stresses that may vary along the cell length (unlike cells stretched solely from their ends), it would not be surprising that the mechanical responses of isolated cells and cells in tissue might differ. Although the study of isolated cells is appealing because of the precise experimental control that is available, one limitation of examining the process of MEF is a sampling problem. Single cell experiments in several labs have shown for example, significant alterations in action potential with acute axial stretch (Riemer and Tung, 2003). However, the numbers of cells in each study are typically small (< 10), and might not be adequately representative of the intact tissue. In one exception, experiments on 43 guinea pig ventricular myocytes revealed two distinct populations with different mechanical properties (Cazorla et al., 1997). In another study, experiments on 57 frog ventricular myocytes showed only a small number to be mechanosensitive, but these cells exhibited robust mechanical responses (Riemer and Tung, 2003). Hypersensitive cells need not be present in great numbers in order to confer mechanosensitivity to cardiac tissue, much like stretch receptors in skeletal muscle. Thus, it may be essential to examine the behavior of tens or hundreds of thousands of cells, which is possible in cultured cell monolayers. Yet another advantage of cell culture models is the potential to study cumulative changes in cellular physiology with prolonged exposure to a given mechanical perturbation, as has been the case with vascular endothelial cells (Sato and Ohashi, 2005), vascular smooth muscle cells (Stegemann et al., 2005), chondrocytes (Smith et al., 2004), pulmonary epithelial cells (Trepat et al., 2006), pulmonary endothelial cells (Birukova et al., 2006), cortical neurons (Cohen et al., 2007) and other cell types.

Presently available cardiac cell culture models are derived from neonatal rat, neonatal mouse or embryonic chick, although new cell models are becoming available from human stem cell sources (Gepstein, 2002). Confluent monolayers of cultured cardiac cells constitute a minimal level of complexity that still captures many of the salient features of intact tissue function, yet they are simple enough that the tissue parameters can be controlled systematically. This intermediate level of complexity between cell and tissue provides a quantitative biophysical model that is experimentally accessible to reproducible conditions such as different forms of mechanical perturbations. Unlike conventional tissue preparations, the two dimensional (2-D) monolayer structure can be controlled using methods of tissue engineering to direct the pattern of cell growth (via engineered surfaces and substrates) (Rohr et al., 1991; Bursac et al., 2002) and the composition of cell types (Abraham et al., 2005; Chang et al., 2006), so that potentially, the mechanosensitivity of different cell types or mixtures can be characterized. Because the 2-D cell culture allows rapid superfusion and control of the extracellular environment, they are well suited for detailed biophysical experiments, especially when applying different drugs or molecular interventions to identify signaling pathways of interest in MEF.

2.2 MEF in NRVM monolayers

A wide variety of laboratory apparatuses can be used for mechanical stimulation of cell and tissue cultures, including compression, longitudinal stretch, bending, axisymmetric substrate bulge, in-plane substrate distention, fluid shear stress, and combined substrate distention and fluid shear stress (Brown, 2000), and also magnetic twisting and pulling cytometry (Lele et al., 2007). The following studies are specifically those that applied stretch or shear stress to cardiac cell monolayers and provide clear evidence that MEF is involved in different signaling pathways in cardiac cell functions and may play an important role in cardiac diseases caused by increased mechanical loading of the ventricular myocardium, such as dilated cardiomyopathy or hypertrophy. For a general review of myocardial stretch effects at the molecular and cellular levels as they relate to hypertrophy in NRVMs, see Sadoshima and Izumo (1997).

Several groups have shown that cyclic stretch upregulates the expression levels of Cx43 in cultured monolayers of neonatal rat ventricular myocytes (NRVMs) grown on silicone membranes (Wang et al., 2000; Zhuang et al., 2000; Shyu et al., 2001; Pimental et al., 2002; Shanker et al., 2005; Yamada et al., 2005). The stretch-induced upregulation of Cx43 has been related to activation of the Na-H exchanger (Wang et al., 2000), increased secretion of angiotensin II that acts on AT1 receptors (Shyu et al., 2001), secretion of vascular endothelial growth factor (VEGF) acting downstream of TGFβ and FAK in an autocrine fashion (Pimental et al., 2002; Yamada et al., 2005), and type of extracellular matrix protein on which the cells are grown (Shanker et al., 2005). Uniaxial cyclic stretch induces cell alignment of NRVMs (Terracio et al., 1988; Vandenburgh et al., 1995; Matsuda et al., 2005) and localization of Cx43 at the longitudinal cell termini that is regulated by the Rac1 pathway downstream of N-cadherin (Matsuda et al., 2006). Anisotropic biaxial static stretch of micropatterned NRVMs grown as linear strands showed upregulation of Cx43 and N-cadherin with transverse stretch but not with longitudinal stretch (Gopalan et al., 2003). In other studies, cyclic stretch of NRVMs grown on silicone membranes coated with RGD or YIGSR peptides were similar in terms of cell adhesion to NRVMs grown on fibronectin or laminin, respectively, but not in terms of normal FAK expression or sarcomere formation (Boeteng et al., 2005). In the experiments described above, stretch was achieved either with vacuum pulses applied to a Flexcell strain unit (Flexcell International Corp., Hillsborough, NC) (Wang et al., 2000; Shyu et al., 2001; Boateng et al., 2005) or with experimental chambers that were custom designed to allow uniaxial stretch (Terracio et al., 1988; Zhuang et al., 2000; Matsuda et al., 2005) or biaxial stretch (Gopalan et al., 2003).

Electrophysiological measurements have been successfully carried out using optical mapping of NRVM confluent monolayers on thin silicone membranes (Zhuang et al., 2000). Propagation velocity increased from 27 to 35 cm/s after 1 hr pulsatile stretch and to 37 cm/s after 6 hrs pulsatile stretch. Upregulation of proteins that form electrical and mechanical junctions was also observed, although with no significant change in the upstroke velocity of the action potential or cell size. Static stretch produced qualitatively similar but significantly smaller changes than pulsatile stretch. However, these responses require an hour or more and involve transcriptional and post-translational mechanisms. There have been fewer studies of immediate acute responses to mechanical loading of myocytes in cell culture.

NRVM monolayers continuously exposed to low fluid shear rates (5–50/s) showed an immediate, graded and reversible increase in their spontaneous beating rate (up to 500%) that was substantially attenuated either in the presence of isoproterenol or incubation with integrin-blocking RGD peptides (Lorenzen-Schmidt et al., 2006). These findings suggest that the β-adrenergic signaling pathway and integrin activation are involved in the shear stress response.

Optical mapping experiments in which pulsatile jets of fluid were impinged upon sheets of cardiac cells showed that the jets are capable of exciting monolayers of cardiac cells, causing propagated action potentials and induce reentry (Kong et al., 2005). A fluid jet impinging on a planar surface consists of both compressive and shear stress elements. Applying compressive force (pressure) and shear stress separately to monolayers of cardiac myocytes may help to better understand their individual roles in MEF in hearts, as presented in the next section.

3. Experimental methods to apply hydrostatic pressure, fluid shear stress and anisotropic stretch to NRVM monolayers

3.1 NRVM monolayer culture and optical mapping system

Ventricular myocytes were obtained by enzymatic dissociation of minced ventricle tissue obtained from 2-day old Sprague-Dawley rat pups (Bursac et al., 1999), Following two preplating steps to lower fibroblast concentration, myocytes were then plated onto fibronectin-coated coverslips at a concentration of 0.5 million myocytes per cm2 and cultured under 95%O2/5% CO2 at 37 °C in a incubator for 5~6 days before experiments. The experimental Tyrode’s bath solution consisted of (in mM): 135 NaCl, 5.4 KCl, 1.8 CaCl2, 1 MgCl2, 0.33 NaH2PO4, 5 HEPES, and 5 glucose.

The temperature of the cell monolayer was kept at 34~36°C by flowing warm Tyrode’s solution into the chamber in between experimental runs and keeping the experimental chamber within a heated enclosure to maintain the temperature. Experiments were carried out inside an enclosed chamber with controlled temperature in the range of 30~36°C.

The optical mapping system used in the experiments has been previously described (Entcheva et al., 2000; Lim et al., 2006). The cell monolayer was stained for 5 min in the dark with 10 μM of the voltage-sensitive dye di-4-ANEPPS (Molecular Probes). Optical action potentials were recorded from 253 sites with 1mm spatial resolution. Each optical recording was taken for a period of 2 sec with 1000Hz data sampling rate, from which maps of the activation patterns of the cell monolayer could be obtained.

3.2 Application of hydrostatic pressure

To apply a uniform pressure with no shear, the monolayer was placed in a sealed chamber under no-flow conditions. The static pressure chamber is shown in Fig. 1A. A mechanical switch was connected to the pressure port, and could be turned on and off manually with controlled timing.

Fig. 1.

Fig. 1

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Experimental chambers used in experiments. Panel A. Closed chamber for pressure. Panel B. Parallel plate flow chamber for shear stress. Panel C. Anisotropic stretching device.

Point stimulation was applied with frequency of 2Hz, 3Hz and 4Hz. The pacing pulses were 10 ms biphasic rectangular pulses with amplitude 1.5X diastolic excitation threshold. The applied pressure was either atmosphere pressure (Patm) or 110mmHg above Patm. Each pressure was applied for about 5 min, after which three recordings were taken under that pressure condition about 1 min apart. Afterwards, the pressure was decreased back to atmospheric pressure, and fresh warmed Tyrode’s solution was superfused for 2 min, after which another experimental cycle with different pacing frequencies would be repeated.

3.3 Application of shear stress

Uniform and steady shear stress across a cultured cell monolayer can be generated experimentally in a parallel plate flow chamber (Bakker et al., 2003). To generate laminar flow on top of the cell monolayer, the height of the chamber should be much less than the overall chamber width, and the entrance and exit width should be much smaller than the chamber length. The flow chamber size that we used was 10cm x 3.8cm x 0.05cm (Length x Width x Height). The shear stress τ on the monolayer surface is uniform over the central region of the chamber and is determined by the formula, τ = 6Qμ (wh2), where Q is the volume flow rate, μ is the absolute fluid viscosity, and w and h are the width and gap height of the rectangular flow channel, respectively. With a solution flow rate of 0.23 ml/s, a shear stress of 1 dyn/cm2 could be generated.

The flow chamber is shown in Fig. 1B and consists of a milled polycarbonate cover plate, a thin rectangular silicon rubber gasket, and a milled polycarbonate bottom plate. The coverslip with the attached cell monolayer was placed within the bottom plate which contains a milled well with a depth and diameter that matched the thickness and diameter of the slide, so that the cells were exposed to an uninterrupted laminar flow.

Cells were stimulated through a pair of line electrodes by 10 ms biphasic rectangular pulses with amplitude 1.5X diastolic excitation threshold. For short exposure to shear stress experiments, zero (no flow) and 1.0 dyn/cm2 shear stress were applied. The zero shear stress condition was applied for about 3 min, and two recordings were then taken. After turning on the shear stress, another recording was taken within 1 min. Cells were constantly paced at 3Hz throughout the experiments.

For longer exposure experiments to shear stress, shear stresses of 0.19 dyn/cm2 and 1.1 dyn/cm2 were used. Each shear stress condition was applied for about 4~6 min, and three recordings were taken about 1~2 min apart. The APD80 from all analyzable channels and CV values along 3 parallel paths in the low shear stress condition were used as reference; thus they were zero and had no error. For all later recordings, APD80s and CVs were compared with the baseline values, channel by channel for APD80 and path by path for CV. Cells were constantly paced at 2Hz throughout the experiments.

3.4 Application of anisotropic stretch

An anisotropic stretch device developed for NRVM cell cultures has been previously described (Gopalan et al., 2003) and used in these experiments, and is shown in Fig. 1C. The cells were cultured on an elliptically-shaped, deformable elastomer (shadowed area in Fig. 2A), that was clamped into an elliptical groove in the bottom of the membrane holder by a rubber O-ring. A polycarbonate indenter was fit inside the membrane holder and pushed down by a flange when the screw top was turned. By twisting the screw top, the elastomer and attached cell monolayer could be stretched anisotropically with a 2:1 ratio, as determined by the eccentricity of the indenter. Strains of 10%:5% were used in the experiments. The direction of maximum principal stretch was along the minor axis of the ellipse (shown in Fig. 2A). The strain pattern of the membrane was homogeneous (except within 2 mm of the indenter ring) and anisotropic, and its magnitude could be controlled by the depth of the indenter that was pushed down by the screw top. An additional plastic plate was machined to insert into the top open space to position field electrodes about 2~3mm above the cell monolayer. A mechanical adaptor was also designed to hold the stretch device in place relative to the optical mapping array, so that the mapping locations could be determined with an accuracy of about 0.5mm before and after stretch. Fig. 2 shows the experimental conditions used for anisotropic stretch.

Fig. 2.

Fig. 2

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Schematic of anisotropic stretch experiment. Panel A. Isotropic cell culture consisted of 5 million cells plated in the elliptical gray area. Dark circle indicates area 17mm in diameter scanned by the optical mapping array. Electrical field stimulation consisting of biphasic pulsees was applied through line anode and cathode at opposite sides of the mapping area. Anisotropic stretch with aspect ratio of 2:1 was applied as shown. Layout of 253 optical recording channels is shown in panel B. Each channel records electrical activities from hundreds of cardiac cells inside a circular area 1 mm in diameter.

Cells were stimulated by a pair of line electrodes by 10 ms biphasic rectangular pulses with amplitude twice diastolic excitation threshold. Three recordings were taken 1 min apart prior to stretch. After stretch was applied, 3 more recordings were obtained 1 min apart within the next 10 min. Cells were paced at 2 Hz throughout the experiments.

3.5 Data analysis

Recording sites that had unstable baselines or excessive noise were not included in the analysis. Action potential duration at 80% repolarization (APD80) and conduction velocity (CV) were measured under hydrostatic pressure, shear stress or anisotropic stretch conditions and compared with their corresponding values under control conditions. For each mechanical intervention, APD80s (and CVs) for each site (and path) were averaged over all of the beats occurring during each 2-sec recording period, and then their means were grouped together and compared with their values similarly obtained under control conditions. Data was analyzed with the two-tailed paired Student’s t-test and F-test. P<0.05 was treated as statistically significant.

4. Experimental results

4.1 NRVM monolayer under uniform compressive force (hydrostatic pressure)

APD80s and CVs under two different pressures are shown in Fig. 3. APD80 values drop with increased pacing rates, whereas at a given pacing rate, their changes with pressure were not statistically significant. CV was measured along 3 defined paths for each cover slip, and overall was not found to be significantly changed.

Fig. 3.

Fig. 3

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Acute pressure effects for NRVM monolayer paced at different frequencies. Optical recordings were taken at atmospheric pressure and 3~5 min after high (110 mm Hg) pressure was applied. Panel A. Action potential durations measured at 80% repolarization (APD80), averaged over all recording sites (P>0.05 for all pacing rates, n=1596 recording sites in 7 monolayers). Panel B. Conduction velocities (CVs) averaged over 3 paths in each monoylayer (P>0.05 for all pacing rates, n=21 paths in 7 monolayers).

4.2 NRVM monolayer under shear stress

For the short duration experiments, action potentials measured from all of the recording sites were compared either between successive sets of paced beats under control conditions, or between sets of paced beats before and after ~1 min application of shear stress. A sample result of the changes in action potential duration at 80% repolarization (ΔAPD80) is shown in Fig. 4.

Fig. 4.

Fig. 4

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Change in action potential duration (ΔAPD80) following application of 1.0 dyn/cm2 shear stress for the different recording sites in the monolayer. A small number of sites showed changes exceeding ±15%, as illustrated for sites marked by the oval and square. Individual traces are plotted in panel B with ΔAPD80 > 15% and panel C with ΔAPD80 < –15%.

Three groups of APD measurements were taken from a total of 1501 sites distributed over 9 monolayers: group 1 (control conditions, from a single average of 6 successive beats), group 2 (control conditions, from an average of 6 successive beats taken ~1 min after group 1) and group 3 (fluid shear stress conditions, from an average of 6 successive beats taken ~1 min after group 2). ΔAPD80 decreased from 151.8±23.1 ms to 150.2±24.2 ms to 148.2±23.3 ms, respectively, and did not reach significance for group 2 vs. group 1 (P = 0.14) but was significant for group 3 vs. group 2 (P < 0.001). The scatter in ΔAPD80 levels followed a Gaussian distribution that was observed in all three groups and could not be attributed directly to the application of shear. F-test analysis of the variance in APD values did not reveal a significant change for shear stress compared with control. CV measured along 3 defined paths for each monolayer (n=21 paths from 7 monolayers), was 21.2±2.1 cm/s for control vs. 20.5±1.7 cm/s with fluid shear stress. However, the decrease in CV was not statistically significant (P = 0.31, n=21).

For the experiments involving longer exposure to fluid shear, low and high shear stress (0.19 dyn/cm2 and 1.1 dyn/cm2 ) were applied twice, each time with a duration of 4~6 min, in order to check the recovery of the electrophysiological functions. Experimental results on APD80 and CV are shown in Fig. 5. APD80 increased significantly each time the shear stress was high, and decreased each time the shear stress was low (P<0.05, n=238 recording sites). For APD80s at the same level of shear stress, no statistically significant differences were observed. CV decreased significantly each time the shear stress was high, and increased each time the shear stress was low (P<0.05, n=3). For CVs at the same level of shear stress, no statistically significant differences were observed (P>0.05, n=3). Qualitatively similar results were observed in 9 other monolayers.

Fig. 5.

Fig. 5

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Repeated APD80 changes (panel A) and CV changes (panel B) averaged over all recording sites when a cell monolayer was exposed to 0.19 dyn/cm2 and 1.1 dyn/cm2 shear stress. * indicates statistically significant changes (P<0.05, n=238 recording sites for APD80, n=3 for CV) compared with level just prior to elevation of shear stress.

4.3 NRVM monolayer under anisotropic stretch

Five cell monolayers were tested for the stretch experiments. All had similar results — APD80 was prolonged, and CV was slowed after stretch. One pair of typical action potential (AP) traces that prolonged by stretch is plotted in Fig. 6A. Fig. 6B shows averaged APD80 changes for all analyzable channels in the five monolayers before and after stretch. Fig. 6C shows the distribution of the change in APD80s in all five monolayers, which increased in a majority of the cells. CV changes are shown in Fig. 7. CV for the 3 paths in one monolayer was 28.1±1.1 cm/s before stretch (Fig. 7A), and 26±1.6 cm/s after stretch (Fig. 7B), about a 7.5% reduction. Fig. 7C compares CVs before and after stretch for all of the selected paths in five monolayers. Both APD80 prolongation (P<0.001, n=503 recording sites) and CV slowing were statistically significant (P=0.037, n=9).

Fig. 6.

Fig. 6

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Panel A. Typical normalized AP traces before and after anisotropic stretch of 10%:5% (paced at 2Hz). Panel B. APD80 before and after stretch. Each recording site in each monolayer was assumed to be is statistically independent from the others with regard to stretch effects, and hence were aggregated into a composite group. Changes are statistically significant (P<0.001, n=503 recording sites). Panel C. Histogram of changes in APD80 with stretch. APD80 was prolonged among the majority of cells (total of 503 recording sites across 5 monolayers). Mean is indicated by the dashed white line and was 6.9%.

Fig. 7.

Fig. 7

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Isochrone maps before (panel A) and after (panel B) anisotropic stretch of 10%:5%. 5% stretch was applied along the direction of wave propagation (the vertical direction). The location of each optical recording channel is shown (+ symbol). Color bars and labels indicate the activation time (in ms) of each recording channel. Time interval between isochrones is 3 ms. Paths chosen for the 3 CV measurements are shown by the straight lines. Panel C. Overall CV changes in all 5 monolayers are statistically significant (P=0.037, n=9). CV for each path in each monolayer was assumed to be statistically independent from the others with regard to stretch effects, and hence were aggregated into a composite group.

5. Discussion

Mechanoelectric feedback (MEF) alters the heart’s electrical activity through changes in the cardiac mechanical environment. Previous work in the literatures on single cardiac cells and whole hearts has shown that mechanical stretch can alter action potential duration (Zeng et al., 2000), and generate extrasystoles in isolated frog (Lab, 1978), pig (Dean and Lab, 1989) and dog (Hansen et al., 1990; Stacy et al., 1992) hearts. It has been speculated that this process of “mechanoelectric feedback” could alter the normal distribution of repolarization and excitability (Tung and Zou, 1995; White et al., 1993), and potentiate the likelihood for reentrant arrhythmia in failing hearts (Dean and Lab, 1989; Reiter, 1996). We are testing MEF effect experimentally using 2D cardiac cell monolayers, characterizing electrophysiological properties, such as excitability, refractoriness (action potential duration) and electrical conduction (conduction velocity) of the monolayer under different mechanical conditions. The employed mechanical forms include hydrostatic pressure, shear stress and stretch, or their combinations. The ultimate goal is to study the mechanisms of MEF inside the hearts and illustrate how different mechanical forms facilitate the initiation of arrhythmic activity.

Within the applied pressure range, no statistical significant changes were observed. This result is not unexpected, considering that pressure was applied uniformly around the cardiac cells. No pressure gradient exists inside and outside the cell membranes or among the cardiac cells. This experiment was attempted to confirm the lack of electrophysiological changes when cardiac cells are subjected to elevated pressure conditions. For the time intervals that we tested, changes in electrophysiological function were not seen. It would be interesting to determine what might happen if cells were cultured long term under abnormal pressure conditions.

For shear stress experiments, no statistically significant changes were observed for brief (<1 min) application of shear stress (up to 1.0 dyn/cm2), although APD80 prolongation and CV slowing were seen for longer intervals (>1 min) of shear stress application (up to 1.1 dyn/cm2). It is believed that shear stress influences cell function through different mechanosensitive structures in the membrane, such as adhesion receptors and associated signaling complexes (Janmey and McCulloch, 2007). Moreover, deformations of cells under shear stress conditions may be subcellular in length scale (Barbee, 2002) and may be fundamentally different in nature than those involving axial stretch. The time delay in shear stress effects might be related to the mechanical properties of cytoskeletal networks which can exhibit viscous characteristics (Cooper, 2006).

For short term exposure to fluid shear stress, although statistically significant changes in APD80 were not observed, certain sites (outliers) in the monolayer exhibited particularly prolonged or shorten APD80 (Fig. 4). These outliers might be indicative of mechanically hypersensitive cells, as suggested previously (Riemer and Tung, 2003). Optical mapping at a microscopic scale can make possible the simultaneous measurement of transmembrane potential from a large number of recording sites with spatial resolution down to 4 μm (Rohr and Kucera, 1998). The ensuing high spatial resolution propagation maps generated from such recordings can potentially identify the size and distribution of local “hotspots” of mechanically perturbed action potentials and propagated activity.

Finally, the prolongation in APD (Fig. 6) and slowing of CV (Fig. 7) with anisotropic stretch are novel and consistent with whole heart observations (Sung et al., 2003). These new observations suggest for the first time that load-dependent conduction slowing in the intact ventricle is likely to be an intrinsic multicellular response (possibly mediated by stretch regulation of cell-cell junctions) rather than a secondary phenomenon related to alterations in tissue perfusion, interstitial flow or extracellular resistances. Future studies into the mechanism for these changes are possible with the cell culture model. Drugs such streptomycin and gadolinium that block certain stretch-activated channels may provide additional insight into the basis for the stretch responses.

6. Conclusion

Mechanically-induced alterations in the electrical behavior of the heart may have serious pathophysiological and clinical consequences. Their better understanding is important for guiding the development of effective treatment strategies. By applying compressive force (hydrostatic pressure), shear stress and anisotropic stretch to cardiac monolayers, quantitative measurements of electrophysiological functions (action potential and conduction velocity) at multicellular level can be obtained. The contemporary cardiac cell monolayer model is a valuable tool that spans the gap between the single cell and multicellular tissue level to aid in the study of mechanical forces and biophysical mechanisms involved in MEF.

Acknowledgments

Funding for this work was provided by an AHA Mid-Atlantic Affiliate postdoctoral fellowship (to Y. Z.), National Institutes of Health grants R01 HL66239 (to L.T.), R21 RR017073 (to L.T.), P01 HL46345 (to A.D.M.) and NSF grant BES-0506252 (to A.D.M.).

Footnotes

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