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Am J Physiol Heart Circ Physiol 285: H1849-H1856, 2003. First published July 10, 2003; doi:10.1152/ajpheart.00384.2002
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Pressure overload alters stress-strain properties of the developing chick heart

Christine E. Miller,1 Chandra L. Wong,1 and David Sedmera2

1Division of Pediatric Cardiology, University of Rochester School of Medicine and Dentistry, Rochester, New York 14642; and 2Department of Cell Biology and Anatomy, Medical University of South Carolina, Charleston, South Carolina 29425

Submitted 9 May 2002 ; accepted in final form 2 July 2003


   ABSTRACT
TOP
 ABSTRACT
MATERIALS AND METHODS
 RESULTS
 DISCUSSION
 DISCLOSURES
 REFERENCES
 
As a first step in investigating a control mechanism regulating stress and/or strain in the embryonic heart, this study tests the hypothesis that passive mechanical properties of left ventricular (LV) embryonic myocardium change with chronically increased pressure during the chamber septation period. Conotruncal banding (CTB) created ventricular pressure overload in chicks from Hamburger-Hamilton (HH) stage 21 (HH21) to HH27, HH29, or HH31. LV sections were cyclically stretched while biaxial strains and force were measured. Wall architecture was assessed with scanning electron microscopy. In controls, porosity-adjusted stress-strain relations decreased significantly from HH27 to HH31. CTB at HH21 resulted in significantly stiffer stress-strain relations by HH27, with larger increases at HH29 and HH31, and nearly constant wall thickness. Strain patterns, hysteresis, and loading-curve convergence showed few differences after CTB. Trabecular extent decreased with age, but neither extent nor porosity changed significantly after CTB. The stiffened stress-strain relations and constant wall thickness suggest that mechanical load may play a regulatory role in this response.

cardiac development; trabeculae; left ventricle

MANY CELL TYPES SENSE AND respond to mechanical stimuli, producing biochemical signals that control gene expression, in turn affecting tissue mass, shape, and/or physical properties. It is well known that hemodynamic pressure overload in the mature left ventricle (LV) results in increased mass with reversible or irreversible LV hypertrophy (1, 9, 18, 34). Whether physical properties change, in particular passive elastic stress-strain relations, is questionable; similar (37, 41) and significantly steeper (11, 36) diastolic stress-strain relations have been observed after aortic banding-induced pressure overload in adult mammals.

Embryonic cardiac muscle also responds to increased hemodynamic pressure with increased mass, but by cardiomyocyte hyperplasia instead of hypertrophy (8). Morphological and myoarchitectural changes occur, including chamber dilatation, double-outlet right ventricle (RV), persistent truncus arteriosus, ventricular septal defect, thickening of the compact myocardium and trabeculae, and spiraling of the trabecular course (8, 40). Similar and reduced LV stress-strain relations have been reported after ventricular pressure overload (44).

Understanding the role of mechanical factors in heart development is important because of the critical function of the organ and the potential for congenital defects. Several studies suggest that mechanical stress and strain may play an important role. A cardiac model incorporating end-diastolic sarcomere length and early-systolic stretch in feedback showed physiological adaptations (2). Stress-related growth caused by end-diastolic pressure in a model of the chick ventricle at Hamburger-Hamilton (HH) stages 21–29 correlated well with experimental results (28). Cultured myocardial cells from HH31 chick were shown to proliferate in response to mechanical strain in vitro (31). A fundamental question is whether a negative-feedback control mechanism exists, so that overall tissue response reduces the difference between a mechanical quantity sensed by the cell and a preset limit. The passive mechanical structure of the heart, consisting of geometric and material properties, determines the magnitude of stress and strain for any given hemodynamic pressure. Thus changes in passive mechanical structure are a likely effector in such a feedback scheme.

We hypothesized that the passive material properties of developing LV myocardium change in response to chronically increased LV pressure. To test this hypothesis, LV pressure overload was created by conotruncal banding (CTB) in the embryonic chick heart. The chick heart parallels the human heart in development, has been used extensively in studies of embryonic cardiac function (6, 20, 24), and allows surgical interventions with reincubation to later stages. CTB was applied at HH21 (19), and hearts were studied at HH27 (embryonic day 5), HH29 (embryonic day 6), and HH31 (embryonic day 7). These stages encompass the crucial period of chamber morphogenesis. Stress-strain relations were constructed from uniaxial cyclic loading with biaxial surface strain recording of excised, passive LV sections from normal and pressure-overloaded hearts. To eliminate possible bias from loading direction, sections were excised and loaded in two approximately perpendicular orientations. Wall thickness; cyclic energy loss (hysteresis); normal, transverse, and shear strains; and stress-strain curves were constructed for each individual LV section. Scanning electron micrographs (SEM) from normal and overloaded hearts were also examined to measure the extent of the compact and trabecular myocardial regions and the percentage of intertrabecular spaces. These results were used to adjust the measured stress magnitudes for trabecular porosity.


   MATERIALS AND METHODS
TOP
 ABSTRACT
 MATERIALS AND METHODS
RESULTS
 DISCUSSION
 DISCLOSURES
 REFERENCES
 
Fertilized White Leghorn chicken eggs (Truslow Farms, Chestertown, MD, and American Selected Products, Milton, PA) were incubated blunt end upward at 38.5°C and 62% humidity in a forced-downdraft incubator to the desired stage. CTB was performed at HH21 (embryonic day 3.5), which corresponds to embryonic day 10 in the mouse and Streeter Horizon stage XIII in humans (42). The egg was positioned under a dissecting microscope, and access to the embryo was obtained by opening the shell and incising a small region of the outer and inner shell membranes. A loop of 10-0 nylon suture was tied in a secure but nonbinding manner around the outflow track of the developing heart. Embryos with visual signs of bleeding were discarded. The opening in the shell was sealed with Parafilm, and the egg was reincubated until sampling time. The shells of the matched controls were left intact, and the matched controls were not subjected to the banding procedure.

The CTB procedure increases LV pressure acutely and chronically, with no changes in heart rate or cardiac output (7, 25). At HH21, the interventricular septum (IVS) is not formed; the ventral and dorsal endocardial cushions unite at HH28, and the interventricular foramen closes at HH34 (embryonic day 8). Also, the conotruncus has not septated into aorta and pulmonary artery, which occurs proximally at HH34 (42). Thus the CTB procedure increases afterload to the primitive RV and LV, although only the LV is studied here.

Mechanical testing was performed on tissue from HH27, HH29, and HH31 (embryonic days 5, 6, and 7, respectively) hearts that had undergone CTB at HH21 and on matched intact-shell controls. For all stages tested, the extraembryonic splanchnopleure and somatopleure were cut along the border of the area opaca for removal of the embryos from the yolk. The excised embryo was placed in a petri dish of oxygenated 37°C Krebs-Henseleit buffer with 4 x 10–4 mmol/l verapamil, 30 mmol/l KCl, and 10 mmol/l EGTA. These additives inhibit Ca2+ influx into myocardial cells through slow Ca2+ channels, raise cellular resting potential to inactivate Na+ channels, and chelate intracellular Ca2+, respectively, with the total effect being to prevent Ca2+ release from the sarcoplasmic reticulum and, thus, prevent activation of actin-myosin cross bridges. Myocardial cells are thus maintained in a passive state. A major part of the damage in cutting through myocardial cell membranes is due to the destructive effects of the resulting contracture and supercontracture. Cells interior to the cut can be affected, because the extreme shortening causes adjacent cells to be torn open; hence, the initial cutting injury at the surface is propagated throughout the rest of the tissue by electrical and mechanical propagation (35). Thus the verapamil, KCl, and EGTA additives, which inhibit Ca2+ currents and prevent cross-bridge formation, minimize the dissection injury.

Cutting through the atrioventricular and bulboventricular grooves removed the LV and RV. The ventral and dorsal halves of the ventricles were then separated. Hexahedral sections were cut with microsurgical scissors from the ventral half of the LV (Fig. 1A). Longitudinal (L) sections were oriented in an apicobasal direction, i.e., longitudinal from the apex to the mitral orifice, approximately parallel to the IVS. Section width was approximately equal to section thickness, and total length varied from 400 µm (HH27) to 750 µm (HH31). Transverse (T) sections were approximately perpendicular to L sections and parallel to the outer curvature of the heart; width and length were as described for L sections (Fig. 1B). To HH29, the main trabecular sheets in the LV are dorsoventrally oriented and perpendicular to the outer compact layer. In cross section, they appear parallel to the lateral wall of the IVS, with very fine trabecular linkage to the compact layer (39). The L sections are therefore parallel and perpendicular to primary trabecular orientation at these stages. At HH31, the predominant orientation of the main LV trabecular sheets has changed from dorsoventral to apicobasal as they start the process of fusion to form the papillary muscles. Thus, at HH31, T sections are parallel and L sections are perpendicular to primary trabecular orientation. Images of all excised sections were recorded on videotape before testing.



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  		Fig. 1. A: stereomicroscopic view of a chick heart at Hamburger-Hamilton (HH) stage 29 (HH29) showing approximate location of excised sections. L, longitudinal section; T, transverse section; LV, left ventricle; RV, right ventricle; LA, left atrium; RA, right atrium; IVS, interventricular septum; CT, conotruncus. B: schematic diagram showing excised hexahedral LV section from HH29 chick heart. Dimensions are means of all tested samples.

 

The excised sections were placed onto two 0.38-mm-diameter, stainless steel, bevel-tipped wires coated with {alpha}-cyanoacrylate adhesive and adjusted to a separation ~50% of the length of the ventricular section. The wires extended into a water bath filled with the warmed, oxygenated Krebs-Henseleit buffer and connected to an ultrasensitive force transducer and a linear actuator, as previously described (32, 33). At least three 15-µm spheres of styrene divinylbenzene were placed on the anterior (myocardial) surface of the mounted ventricular section midway between the ends to define a central gage length. Custom LabVIEW software programs (National Instruments, Houston, TX) with an analog-to-digital–digital-to-analog board (model AT-MIO-16XE, National Instruments) controlled the motion of the linear actuator, reversing direction at preset upper and lower limits, and sampled the transducer signal at 500 Hz. The ventricular section was stretched in a cyclic loading-and-unloading pattern at 0.5 Hz for 10 cycles to 15% longitudinal strain. The x-y positions of the markers were determined during stretching as previously described (32) from digitized images from a charge-coupled device video camera attached to the stereomicroscope (ComputerEyes/RT frame-grabber board, Digital Vision, Dedham, MA) and recorded to disk along with transducer output voltage and time. Strains were calculated from the marker positions and the definition of the symmetrical Green-Lagrangian strain tensor (Eij) (32). With z as the long axis of the tissue section, x as the width, and y as the height, the surface strains were Ezz, Exx, and Ezx. Principal strains were also calculated. The calculated strain is homogeneous within the triangle.

The remaining data analysis was done on recorded files of time, voltage, and strain and on videotapes of tested sections. All voltages were referred to the minimum at the end of the third cycle, and force was calculated as transducer output voltage multiplied by calibrated transducer sensitivity, 9.37 V/g. Width (w), length, and thickness (t) of each tested section were measured from videotape (Scion Image program, Scion, Frederick, MD) by averaging measurements at three locations. Thickness was the total distance from the epicardial to the endocardial surface. Rectangularity was assessed by fitting lines through the epicardial and endocardial surfaces from the side view and by measuring the linearity and parallelity of the lines.

Normalized hysteresis area was calculated as the area between the loading and unloading segments divided by the total area under the loading segment in the fourth loading cycle of the stress-strain plots. Hysteresis convergence was the difference between the strain energy stored during loading in the first cycle and that stored during loading in the fourth cycle.

SEM on a parallel sample population provided data on typical cross-sectional characteristics. These data could not be determined noninvasively on tested tissue sections. SEM images of HH27, HH29, and HH31 normal hearts and hearts banded at HH21 (40) were analyzed. Compact layer area proportion (ACL), the percentage of the total cross-sectional solid area occupied by the compact layer, was measured by an unbiased point-counting technique on digitized SEM images using the Analyze 7.1 image analysis system (CNSoftware, London, UK) on a Hewlett-Packard UNIX platform. Trabecular layer area proportion (ATL) is then equal to 1 – ACL. Trabecular gross porosity (PTL) was determined from the point-counting technique as the percentage of the total area of the trabecular layer (myocardial tissue + spaces) occupied by the intertrabecular spaces. We calculated overall porosity (Po), the percentage of the entire cross-sectional area occupied by intertrabecular spaces, as follows: Po = PTLATL/ATL + (1 – PTL)ACL.

Until HH34, linear dimension shrinkage with SEM has been shown to be 52% (39, 40). Independent experiments (39; unpublished observations) were performed comparing trabecular morphology by SEM with confocal images of whole mount phalloidin-stained hearts in the wet state by the method of Germroth et al. (15) and uncut, optically sectioned hearts stained with antimyosin antibody and cleared according to Kolker et al. (27). The trabecular morphology and shape were in complete alignment, showing that critical point drying in SEM results in homogeneous tissue shrinkage. Because of the uniform shrinkage, the dehydrated state of specimens does not influence the values of porosity. Thus ACL and PTL were not adjusted for tissue shrinkage.

To calculate stress, each tissue section was modeled as a laminated bilayer bar of rectangular cross section with width w and thickness t. Both layers contain identical material, one closely packed and the other webbed. The ends are rigidly held and stretched with force F. End effects are neglected. The average longitudinal Lagrangian (2nd Piola-Kirchoff) stress Tzz in the compact layer is then given by the following equation: Tzz = F/[wt(1 – Po)]. Tzz was plotted vs. surface longitudinal strain (Ezz) for each section. Stress-strain curves were fit by various polynomial and exponential functions. The strain energy (SE) density for the loading portion of each curve () was calculated as the area under the stress-strain curves from 0 to 0.15 strain. Secant stiffness, the stress divided by the strain at each data point, was calculated for all experimental data points from 0 to 0.15 strain for each tested section.

The six normal groups are designated by HH stage and orientation (L or T) as 27L, 27T, 29L, 29T, 31L, and 31T. The six banded groups are 27L-B, 27T-B, 29L-B, 29T-B, 31L-B, and 31T-B. Mean ± SD was calculated for each measurement in the 16 groups. Student's t-test gave the significance of differences between the control and the banded group at each stage and orientation and also between L and T orientations at fixed stage and treatment (control and banded). ANOVA with post hoc multiple-comparison Tukey's test compared differences between quantities across stages. Statistical significance level was P < 0.05.


   RESULTS
TOP
 ABSTRACT
 MATERIALS AND METHODS
 RESULTS
DISCUSSION
 DISCLOSURES
 REFERENCES
 
Point-counting analysis of SEM sections from control hearts showed that the percentage of cross-sectional area occupied by the compact layer in control hearts increased significantly with developmental stage from HH27 (26%) to HH31 (48%; Table 1). Banding at all stages produced no significant changes in the percentage of cross-sectional area occupied by the compact layer. Trabecular layer porosity was similar in HH27 and HH31 (Table 1) but declined significantly from HH29 (54%) to HH31 (48%, P < 0.05). Trabecular porosity and overall porosity in corresponding banded and control groups were similar (P = not significant).


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  		Table 1. LV wall characteristics

 

In micrographs, LV dilatation and elongation with the beginnings of trabecular spiraling and thickening of the compact layer were observed in HH27 hearts banded at HH21. At HH29, banding led to further LV elongation, expansion of the trabecular-free lumen in the apical part of the LV with concomitant reduction of the trabecular extent, thickening of the compact layer, and spiraling of the trabeculae counterclockwise from apex to base (Fig. 2).



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  		Fig. 2. Representative scanning electron microscope images of HH29 embryonic chick hearts sectioned in a transverse plane. A: normal heart, midventricular level. Arrow, trabecular sheet. B: midventricular level from heart banded at HH21 (2.5 days earlier); note absence of RV at this level. Arrows, trabecular sheets. C: normal heart, apical level. D: banded heart, apical level; note thickened compact layer, presence of trabeculae-free lumen, and spiraling trabecular course.

 

The total wall thickness of sections from control hearts increased significantly with age from HH27 to HH29 and leveled off at HH31 (Fig. 3). Total wall thicknesses did not increase significantly after CTB (Fig. 3). The mean section widths (µm) as cut for this study were 325 ± 34 (HH27), 305 ± 48 (HH27-B), 349 ± 39 (HH29), 345 ± 28 (HH29-B), 371 ± 34 (HH31), and 388 ± 64 (HH31-B). These widths were approximately equal to the section thickness; thus the excised sections had an almost square cross section. Widths were also fairly uniform, eliminating possible bias in results due to greatly varying cross-sectional shape.



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  		Fig. 3. LV wall thickness of excised sections from chick embryos at HH27, HH29, and HH31; suffix B denotes hearts subjected to pressure overload by conotruncal banding at HH21. Longitudinal and transverse sections are pooled for each stage. Study population is as follows: n = 17 for HH27, n = 16 for HH27-B and HH31, n = 19 for HH29, n = 18 for HH29-B, and n = 20 for HH31-B. Statistical comparisons were made between normal stages and within a stage for normal vs. banded. For controls, significance is as follows: *P < 0.001 vs. HH27. All other comparisons were not significant.

 

During cyclic loading, peak longitudinal Green's strain (Ezz) on the epicardial surface of all 16 groups was 0.16–0.19 and followed a cyclical pattern in phase with force. The transverse epicardial strains (Exx) were negative in all groups and were –0.21 to –0.37 times longitudinal strain, except in 31L sections, where transverse strain was only –0.09 times longitudinal strain. Exx of L and T orientations were similar, as was Exx after banding, except at HH31. Exx in 31T sections was 3.8 times that in 31L sections (P < 0.05), and Exx in 31L-B sections was 2.8 times that in 31L sections (P < 0.05).

The force-strain relations were well defined and repeatable. The loading and unloading branches were separate, showing typical pseudoelastic behavior. In all 12 groups, the loading branch was nonlinear, stiffening with increasing strain. Most curves showed inflection points between 0 and 0.07 strain. A preconditioning phenomenon was present, with the first loading cycle at significantly higher force than the second and subsequent loading cycles (P < 0.01 in all 12 groups). The loading curves converged within 1% after three cycles, resulting in reductions in strain energy of 33–57%, with corresponding L and T orientations and control and banded treatments similar, except at 29T. The reduction was significantly more at 29T (57%) than at 29L (33%, P = 0.02) or 29T-B (33%, P < 0.001).

The mean hysteresis, or energy loss, in a loading-unloading cycle ranged from 16% (27T) to 24% (29L). Hysteresis was similar between L and T orientations and between control and banded groups.

In control hearts, plots of porosity-adjusted mean Lagrangian (2nd Piola-Kirchoff) stress vs. longitudinal Lagrangian (Green's) strain shifted significantly downward with developmental age from HH27 to HH29 for L and T orientations (Fig. 4). Both stress-strain curves had steeper curvature at HH29 than at HH31; 29L was above 31L but 29T was below 31T. The curves were fit well by a cubic polynomial (all R2 > 0.999), but not by simple exponential equations. However, because emphasis could be too easily shifted between the linear, quadratic, and cubic terms of the polynomial and because introduction of a mathematical description of the curves having no unique physiological foundation was deemed unadvisable, these fits were not used for statistical comparisons. Rather, ANOVA of total strain energy showed P < 0.05 among all three curves at L and T orientations. ANOVA of secant stiffness (Fig. 4) also showed significant differences.



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  		Fig. 4. Mean Lagrangian stress adjusted for trabecular porosity vs. longitudinal Green's strain in excised LV sections from control hearts. Numerical labels denote HH stage. A: longitudinal (L) sections. Study population is as follows: n = 8 for 27L, 29L, and 31L. B: transverse (T) sections. Study population is as follows: n = 9 for 27T, n = 11 for 29T, and n = 8 for 31T. For clarity, error bars (SD) are shown only at 0.05, 0.10, and 0.15 strain. All differences between curves are significant at P < 0.05 for strain energy. *P < 0.05 vs. HH27; +P < 0.05 vs. HH29 for secant stiffness.

 

In control hearts, porosity-adjusted mean Lagrangian stress vs. longitudinal Lagrangian strain curves were very similar in magnitude and shape for L and T orientations at HH27 (Fig. 5A). The curve shapes of 29L and 29T were similar, but with 29L above 29T (Fig. 5B; P = not significant). Curve shapes of 31L and 31T were also similar, but with 31T above 31L (Fig. 5C; P < 0.05 above 0.10 strain).



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  		Fig. 5. Mean Lagrangian stress adjusted for trabecular porosity vs. longitudinal Green's strain in control and banded LV sections at HH27 (A), HH29 (B), and HH31 (C). L, longitudinal orientation; T, transverse orientation; B, banded heart. Hearts were banded at HH21. Study populations are as follows: n = 8 for 27L, 27L-B, 27T-B, 29L, 29L-B, 31L, and 31T, n = 9 for 27T, n = 10 for 29T-B, 31L-B, and 31T-B, and n = 11 for 29T. For clarity, error bars are shown only at 0.05, 0.10, and 0.15 strain. All differences between corresponding normal and banded curves are significant at P < 0.05 for strain energy. *P < 0.05 vs. corresponding control group for secant stiffness.

 

Stress-strain curves from HH27, HH29, and HH31 hearts banded at HH21 shifted to larger stresses than control hearts. The shifts were moderate at HH27, with increases of 1.5–2.0 times in secant stiffness and strain energy (P < 0.05 for T; Fig. 5A). At HH29, banded stresses averaged 5.9 (L) to 7.4 (T) times larger than control (P < 0.001 for L and T; Fig. 5B). The difference was also significant at HH31: 5.0 (L) and 2.9(T) times larger than control (P < 0.001; Fig. 5C).


   DISCUSSION
TOP
 ABSTRACT
 MATERIALS AND METHODS
 RESULTS
 DISCUSSION
DISCLOSURES
 REFERENCES
 
This study shows that pressure overload causes significant changes in the passive mechanical material properties of the early embryonic heart. These changes depend on the duration of the overload. In addition, testing of LV tissue from HH27 to HH31 control hearts provides new information on changes in material properties with tissue maturation.

The geometric and material properties measured here in control hearts are not unexpected for this developmental range. Increases in wall thickness and compact layer proportion from HH27 to HH31 are indicative of the general growth and maturation of the tissue. The nonlinear stress-strain curves (13, 16) and preconditioning phenomenon (4, 12, 14, 23, 30, 32) are typical of soft biological tissue. Less expected was the finding that gross porosity of the trabecular layer decreased only slightly from HH27 to HH31, with most of the decrease in overall porosity attributed to the decrease in the extent of the trabecular layer.

The significant decrease in the myocardial stress-strain relation from HH27 to HH29 and HH31 in control hearts was unexpected. It cannot be explained by changes in thickness of the compact and trabecular layers and porosity, because these factors are accounted for in the area normalization factors. Although the averaged overall porosities are means of a parallel study group (total cross-sectional area is measured on each tested specimen), they appear unlikely to bias the results, because the range of porosity adjustments is small relative to the decrease in stress with developmental stage. Further study is necessary to determine the changes in intercellular, subcellular, and/or extracellular matrix (ECM) components and connections responsible for this stress-strain decrease.

Much larger stress-strain magnitudes have been reported at HH27 from uniaxial loading of excised LV strips (44). The data were modeled by the following equation: stress = a[exp(b * strain)], where b for two perpendicular directions is 7.5 and 7.9 and a is a prestress included to achieve "excision length" and is not explicitly given here. If a is in the range of 50 mg/mm2, as shown in pressure-strain graphs, then Tobita et al. (44) report, e.g., 1.1 kPa unadjusted stress at 0.10 strain, in contrast to the much lower stresses measured here (0.20 kPa unadjusted stress at 0.10 strain). The differences between the present results and those reported by Tobita et al. could be due to the prestressing of the tissue, resulting in shifts to higher stress-strain magnitudes, or to the larger strains used, 0.30 compared with 0.15, which may induce tissue changes. Also, the tissue sections tested by Tobita et al. may be from a different LV location, and morphological studies (39) and measurement of intramyocardial pressure (5) suggest that material properties may vary throughout the LV.

The similarity of stress-strain curves and transverse strains in L and T orientations from HH27 to HH29 indicates that loading orientation does not bias the results. It also implies that the relatively isotropic compact layer, on which strains are measured, dominates the overall mechanical response of the LV myocardium in these tests. Anisotropic LV epicardial strain patterns during active contraction have been reported in the chick at HH24, HH27, and HH31 (43), however; thus passive behavior may be different from active behavior. The difference in stress-strain curves and normalized transverse epicardial strains at HH31 suggests that the trabeculae, which have formed into long, thick bundles (trabeculae carneae) attaching to the ventricular wall or IVS along their length (38), influence the mechanical behavior at this stage.

CTB initiated at HH21 induced significant changes in the stress-strain relation by HH27, 1.5 days of overload, with passive stiffness almost doubling. Longer exposure to overload resulted in much larger increases. This period of overload encompasses HH24, when the mitotic activity of myocardial cells is the highest in the entire incubation period (17, 22). Because stress is normalized to porosity-adjusted cross-sectional area, these changes in passive stiffness are independent of the increased growth observed after CTB at HH21, e.g., the increase in cell number by 43% at HH27 and 24% at HH29, the increase in ventricular weight by 67% at HH27 and 86% at HH29 (7), and the visible elongation of the heart (40). Changes in the ECM may be responsible for the pressure overload-increased stiffness, because the characteristics and extent of the ECM are important determinants of diastolic ventricular properties and function (16). Adult animal models showed increased ventricular collagen (3, 11, 45) and constant collagen concentration (26, 37) after pressure overload.

In contrast to the increase in stress-strain magnitude of 1.5–2.0 times (P < 0.05) at HH27 after CTB found here, another study of myocardial material properties after CTB (44) did not measure any significant changes in LV stress-strain relations at HH27 after CTB at HH21 in pressure-strain and excised strip tests. The reason for this difference is unclear but could be due to a different amount of pressure overload, a different location of strips, or larger strains used for testing.

The present results support the possibility of a biological feedback control system in the developing heart, with deformation or strain involved in the sensing and regulatory process. LV wall stress due to internal pressure is directly proportional to pressure magnitude and ventricular radius and inversely proportional to wall thickness. An increase in wall stress due to pressure overload with no change in tissue material properties produces an increase in wall strain. Tissue remodeling causing increased passive stiffness can decrease wall strain, perhaps to a level that is "normal" for that developmental stage. Although the effect is at the tissue level, the mechanisms for such stress- or strain-regulated control likely reside at the cellular level and may involve a number of cell-signaling mechanisms. For example, increased tyrosine phosphorylation has been observed in HH31 chick myocytes strained in vitro on flexible membranes (31), and increased platelet-derived growth factor-like protein has been observed in banded embryonic chick ventricle (21).

The principal limitations of this study arise from the small size and fragility of the embryonic heart. The advantage of the in vitro direct loading method is accurate measurement of tissue geometry and of applied force magnitude and direction vs. in vivo loading, which requires more modeling assumptions and/or simplifications for extraction of stress-strain relations. Prior studies showed that the extraction and attachment method maintains tissue viability away from the fixed ends (32). The uniaxial loading procedure is necessary because of the small tissue size, but loading of sections oriented in perpendicular directions and measuring biaxial surface strains provide multiaxial information. A limitation of the video-tracking method is strain measurement only on the epicardial surface. However, techniques such as magnetic resonance tagging used for noncontact measurement of internal deformations in the mature heart (10, 29) do not have enough resolution for this application. Hand tying the CTB did not introduce any more variation into the results than that already present as a result of natural variation and experimental noise. In five of six groups, the standard deviations of the stresses in banded heart sections are lower than the standard deviations of the corresponding control heart sections. In the single exception, 29T, the banded standard deviations are only 1.09–1.23 times those from control hearts.

In conclusion, this work shows that ventricular pressure overload during the chamber morphogenesis period of early embryonic heart development causes significant stiffening of the stress-strain properties of LV tissue. This stiffening persists 1.5, 2.5, and 3.5 days after banding. The results also suggest a role for mechanical load in regulating cardiac development. The increased passive stiffness may be a useful adaptation, helping distribute and sustain the increased contractile force required by the pressure overload. It may also be detrimental, because increased passive stiffness can contribute to impaired systolic function. At the least, this response at such a crucial developmental stage has the potential to negatively impact further development of the heart and other embryonic systems, with the effects perhaps remaining for the lifetime of the individual. Thus the role of mechanical load in heart development may be very important, and further elucidation of this role may aid in understanding normal cardiac development and processes leading to cardiovascular disease.


   DISCLOSURES
TOP
 ABSTRACT
 MATERIALS AND METHODS
 RESULTS
 DISCUSSION
 DISCLOSURES
REFERENCES
 
This research was supported by National Heart, Lung, and Blood Institute Grant R01-HL-65908, American Heart Association, New York State Affiliate, Grant-in-Aid 9951019T, and National Heart, Lung, and Blood Institute Specialized Center of Research in Pediatric Cardiovascular Disease Grant HL-51498. Experimental work by D. Sedmera was performed at the Institute of Physiology, University of Lausanne (Lausanne, Switzerland).


   ACKNOWLEDGMENTS
 
We thank Franco Ardizzoni for expert technical assistance with SEM and Martina Sedmerova for assistance with point counting. We acknowledge the support of Bucknell University, particularly Weston Griffin and Dr. Keith Buffinton, for developing the linear actuator.


   FOOTNOTES
 

Address for reprint requests and other correspondence: C. E. Miller, Dept. of Mechanical Engr., Dana Bldg., Bucknell University, Lewisburg, PA 17837 (E-mail: cemiller{at}bucknell.edu).

The costs of publication of this article were defrayed in part by the payment of page charges. The article must therefore be hereby marked "advertisement" in accordance with 18 U.S.C. Section 1734 solely to indicate this fact.


   REFERENCES
TOP
 ABSTRACT
 MATERIALS AND METHODS
 RESULTS
 DISCUSSION
 DISCLOSURES
 REFERENCES
 

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