Annals of Biomedical Engineering ( 2009) DOI: 10.1007/s10439-009-9825-8
Changes in the Mechanical Properties and Residual Strain of Elastic Tissue in the Developing Fetal Aorta SARAH M. WELLS1 and E. JANE WALTER2 1 Department of Physics and Atmospheric Science, and School of Biomedical Engineering, Dalhousie University, 5981 University Avenue, Halifax, NS, Canada; and 2Department of Physics and Atmospheric Science, Dalhousie University, 5981 University Avenue, Halifax, NS, Canada
(Received 21 July 2009; accepted 14 October 2009)
Abstract—Formed almost exclusively during development, arterial elastic fibers must function for the lifetime of the animal. We have observed dramatic structural and mechanical changes in aortic elastic tissue during gestational and postnatal development. Elastic tissue was isolated from bovine aortas: (i) during late pregnancy and (ii) in adults. Changes in the relative content of aortic elastic tissue were assessed, as were the viscoelastic properties and residual strains of purified aortic elastic tissue rings. As aortic elastic tissue content increased during development, its circumference and thickness increased—but with circumference rising faster than wall thickness, causing a relative thinning of the elastic tissue. At the same time, elastic tissue stiffness increased while viscoelastic behavior decreased. Much of these changes were concentrated during late gestational development, such that the changes observed during the short span of late gestation examined (~60 days) were similar in magnitude to those occurring over the much longer postnatal period (~1–2 years). Finally, we observed an approximately threefold increase in residual strain in aortic elastic tissue from fetal to adult life, with most of this increase again occurring in late gestation. These results suggest that rapid remodeling, as well as accumulation, of aortic elastic tissue occurs during late gestation. These changes significantly alter both fetal aortic mechanical properties and residual stresses.
d
Keywords—Elastin, Remodeling, Development, Mechanical properties, Aorta, Residual strain.
emax
LIST OF SYMBOLS e10% E* ro eo
Strain at which 10% of failure load is reached Complex viscoelastic modulus Amplitude of stress sinusoid Amplitude of strain sinusoid
Address correspondence to Sarah M. Wells, Department of Physics and Atmospheric Science, and School of Biomedical Engineering, Dalhousie University, 5981 University Avenue, Halifax, NS, Canada. Electronic mail:
[email protected]
r e h t F Ai wi ti At L Dl li
lgauge V rmax
wwet wdry h Ro E
Phase shift between stress and strain sinusoids Stress Strain Opening angle Time Force Initial cross-sectional area Unloaded sample width Unloaded sample thickness Instantaneous cross-sectional area Sample length Actuator displacement, equal to change in sample length once ring fully collapsed Initial unloaded length between MTS posts, equal to the unloaded diameter of the aortic ring Unstretched gauge length of the tissue, equal to one half inner ring circumference Sample volume Ultimate tensile strength, the maximum stress reached before failure Extensibility, strain at which the sample reached rmax Wet weight Dry weight Aortic wall thickness Unloaded aortic outer radius Quasi-static elastic modulus
INTRODUCTION Elastic fibers are the main structural component of tissues that are required to undergo large static or cyclic deformations yet recoil to their original dimensions with minimal energy loss. Characterized by a low stiffness, high extensibility, and high resilience, they are the major component of the aorta, accounting for 2009 Biomedical Engineering Society
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approximately 40–50% of its dry weight in the thoracic region. In the mature aorta, elastic fibers form concentric layers of fenestrated lamellae that bear most of the wall stress under physiological pressures, thus dominating aortic mechanical properties and determining vascular dimensions under these conditions.13,49 Mechanically, elastic fibers are crucial to aortic Windkessel function. The aortic wall rapidly expands during systole, storing elastic energy and minimizing both the rise in arterial pressure during systole and left ventricular workload. The aortic wall then recoils nearly elastically, maintaining blood pressure and flow in diastole, and thereby smoothing the pulsatile output from the heart.4 Despite their vital role in tissue function, aortic elastic fibers are formed almost exclusively during late gestation and the early postnatal period.21,37,42 Elastic fiber synthesis begins early in fetal life with the formation of a scaffold of fibrillin microfibrils.15,21,26 Elastin synthesis starts later, around mid-gestation in most species. During this time, the protein is deposited within and around the microfibrillar scaffold to form the composite elastic fiber.40,42 Elastin synthesis increases to a maximum in the perinatal period,5,12,34 then falls rapidly, leaving little or no elastin synthesis or turnover in adult tissues.11,26,37,42 The confinement of elastic fiber formation to the perinatal period, the lifetime residency of these fibers and their lack of turnover in mature tissues,43,48,51 together make this a critical period for arterial development. Elastic fibers formed early in development must remain intact and functional for the lifetime of the organism. However, the impact of perinatal elastin synthesis during this developmental stage goes beyond load-bearing. These same elastic fibers influence further vascular morphogenesis, regulating the proliferation and organization of vascular smooth muscle cells.33 The importance of perinatal elastin synthesis is highlighted in the ‘‘fetal programming’’ hypothesis of cardiovascular disease. It is believed that adverse conditions during critical periods of fetal growth lead to irreversible changes in tissue structure and function that ‘‘program’’ an individual for increased risk of disease in later life. Epidemiological studies have consistently shown that both systolic blood pressure1,29 and the incidence of cardiovascular disease in adulthood2,25,30,31 are correlated with decreased birthweight. It is hypothesized that, in fetuses whose growth is impaired, synthesis of elastin in the walls of the aorta and large arteries is deficient, leading to permanent changes in their mechanical properties.36 Over a lifetime, these changes precipitate higher blood pressures and cardiovascular disease.37,38 Understanding the normal structural–functional development of aortic elastin during the critical perinatal
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period is therefore fundamental to our understanding of the mechanisms of the fetal origins of cardiovascular disease. While previous studies have looked at biomechanical alterations in the arterial wall during maturation and aging, much less attention has been given to arterial biomechanics during prenatal development—with little or no focus on arterial elastin. A pair of studies characterized developmental changes in full aortic wall structure and their effects on mechanical properties.56,57 Wells et al.57 associated the perinatal accumulation of elastin in the ovine aorta with an increase in the low-stress elastic modulus, reflecting an increased contribution of elastin to wall mechanics at physiological pressures. While the low-stress elastic modulus is directly related to the elastic fiber content, other factors may also contribute, such as developmental changes in changes in the intrinsic mechanical properties of the elastic fibers themselves. Such changes could certainly occur with any developmental changes in fiber composition (microfibrils versus elastin), and/or changes in the structural geometry of these components. While direct structural and mechanical studies are valuable tools to evaluate arterial development, the stress state in a vessel in the absence of applied loads, and the related residual strains, may also provide important quantitative information on remodeling. The unloaded stress state of an artery may develop due to differential growth and/or remodeling across the artery wall. Residual strain can then be characterized by measuring the opening angle of a cut ring sample of the vessel wall. Fung first proposed that this ‘‘zerostress state’’—unaffected by internal or external loads—provides a standard morphological configuration to describe arterial tissue, and that changes in this state provide a sensitive indicator of arterial wall remodeling.16–18,35 To the authors’ knowledge, changes in the residual strain of aortic elastic tissue during development have not been examined. The linked objectives of the current study were thus to determine changes in both the mechanical properties and zero-stress state of isolated elastic tissue (elastic fibers), using as a model the bovine aorta. The bovine aorta is a suitable aortic size for testing. Moreover, bovine fetal cardiovascular structure and function during gestation parallel those in humans and other mammalian species. For instance, the bovine heart begins to beat at 21 d.g. (days gestation)19 and there is the same gradual, late gestational rise in fetal blood pressure.46 We have examined aortic tissue composition to obtain the relative content of elastic fibers vs. water content, and assessed the viscoelastic properties of purified aortic elastic fiber rings using both quasistatic and dynamic stress–strain tests. We have also
Changes in the Mechanical Properties and Residual Strain of Elastic Tissue
examined changes in the opening angle (residual strain) of unloaded aortic elastic tissue rings.
METHODS Tissue Harvest and Sample Preparation All tissue harvest procedures were approved by the Animal Care Committee of Dalhousie University and were conducted in accordance with guidelines approved by the Canadian Council of Animal Care. Bovine thoracic aortas were obtained fresh from slaughter at a local abattoir. Fetal gestational age was estimated from crown–rump length.50 Aortas were collected from adult animals (males or neverpregnant females) and from fetuses (57–100 cm), ranging in age from ~190 to 270 d.g. (full term), thus covering approximately the final trimester of bovine gestation. Vessels were cut proximally just below the aortic arch and distally above the diaphragm. For transport back to the laboratory, tissue was placed in iced Hanks’ physiological saline, with 1% antibiotic–antimycotic (10,000 units/mL penicillin G, 10 mg/mL streptomycin sulfate, and 25 g/mL amphotericin B solubilized in a proprietary citrate buffer (Sigma– Aldrich, Oakville, ON) and either 0.005 g/L trypsin inhibitor (Fluka Biochemicka/Sigma–Aldrich, Oakville, ON) and 0.035%/L phenylmethanesulfonyl fluoride (PMSF). Three tissue rings were collected from each animal: one for each test protocol (composition assessment, opening angle measurement, and mechanical testing). It was important that consistent anatomic landmarks be used for sample collection from each animal since elastin and collagen content vary with longitudinal position.10 The first aortic ring was thus excised from just above the first pair of intercostals arteries on the proximal aorta. This sample was used for determination of elastic fiber and water content. The remaining aortic rings were then excised starting just below the first proximal intercostals. Rings not required for immediate testing were cleaned of adhering fatty tissue, wrapped, and frozen at 84 C. Elastic tissue (i.e., elastin with intact microfibrils) was purified from the aorta as described by Gratzer and Lee.22 To obtain ‘‘purified elastic tissue’’ samples, aortic rings were autoclaved at 121 C for 90 min in fresh dH2O then rinsed three times in dH2O with constant agitation to remove any remaining collagen and cellular debris. Amino acid analysis was performed on treated elastic tissue samples from each age group to confirm the efficacy of the purification.22,53
Compositional Analysis Whole aortic rings for compositional analysis were blotted on all sides with gauze, and then rapidly placed into tared vials to obtain the wet weights. The rings were freeze-dried and then re-weighed. The whole tissue water content, expressed as a percentage, was determined using the wet and dry weights of the whole aortic tissue sample: WWhole Tissue Dry % Water Content ¼ 1 100 WWhole Tissue Wet ð1Þ After weighing, the dried samples were rehydrated overnight in distilled water at 4 C, then autoclaved and rinsed to obtain purified elastic fibers in the rings, as described above. The purified rings were again freeze-dried and re-weighed. The elastic fiber content (expressed as a percentage of whole aortic tissue dry weight) was calculated from the dry weights before and after autoclaving: Elastic Fiber Content ð% Dry WeightÞ WElastic Tissue Dry ¼ 100 WWhole Tissue Dry
ð2Þ
Opening Angle Measurements Opening angle measurements were made on the purified aorta rings as described by Fung and Liu.18 Each ring was opened with a single radial cut with a razor blade and placed in a petri dish filled with enough distilled water to allow the ring to float freely. After 15 min to allow for viscoelastic effects, the opened ring was imaged using a monochrome CCD camera (model 4912-2010, COHU, San Diego, CA) with a zoom lens (model S6X11, Scion Corp., Fredrick, MD) using a custom-written frame grabber program in LabVIEW (National Instruments, Austin, TX). The lumen of each opened ring was traced in Image J (National Institute of Health, USA, public domain) using the freehand drawing tool and the Pathwriter plugin (Wayne Rasband, NIH, Bethesda, MA). The opening angle was taken as the angle between the lines joining the midpoint of the lumen to its two endpoints at the cut18 (Fig. 1), and was calculated from the trace coordinates using a program custom-written under LabVIEW.
Mechanical Tests Mechanical testing was used to assess the quasi-static and dynamic viscoelastic properties of the purified
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radial cut
θ
midpoint FIGURE 1. Measurement of opening angle. When cut, the tissue ring relaxes open from the so-called ‘‘zero external load’’ state to a ‘‘zero-stress’’ state.18 The opening angle h was taken as the angle between the lines joining the mid-point of the inner lumen to the two cut ends of the inner lumen.
elastic fiber tissue. Forced vibration and stress-to-failure tests were carried out on aortic rings from each animal. Unloaded sample dimensions were measured prior to mechanical testing. Images of the purified aortic rings were captured as described above for opening angles. Inner and outer circumferences of aortic rings were measured from the digitized images using image analysis (ImageJ; using the freehand drawing tool and Pathwriter plugin). Wall thickness measurements were taken using digital calipers at four to eight evenly spaced locations around the purified ring and an average wall thickness was calculated for each ring. Mechanical tests were performed using an MTS (Eden Prairie, MN) servo-hydraulic planar biaxial testing device equipped with four actuators and phased stroke/load waveform synthesis.52 For the present studies, elastic tissue rings underwent uniaxial tests using one pair of actuators only. One actuator was fitted with a GSO series 1000 g cantilever load cell (Transducer Techniques, Temecula, CA) and was held stationary throughout the tests. The movement of the other actuator was controlled by an MTS TRAC waveform synthesizer, programmed using MTS software. Grips were attached to the load cell and end of the moving actuator, with a 5 mm vertical post projecting from each. The purified aortic ring was looped loosely around the posts, and the moving actuator was retracted until both posts just came into contact with the inner edges of the ring. At this point, the inner distance between the posts was measured using digital calipers and was taken as the unloaded length of the aortic ring sample, li, approximately equal to the unloaded inner diameter. Throughout the mechanical testing, aortic rings were immersed in a bath of Hanks’ physiological saline solution maintained at 37 ± 1 C by an immersion heater. Time, actuator position, and force data were acquired using a custom-written LabVIEW program running on a computer equipped
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with a data acquisition card (National Instruments PCI-MIO-16E-4). To ensure that the forced vibration tests were nondestructive (especially for the delicate elastic tissue rings from the youngest fetuses), tests were carried out at deformations well below the elastic limit of the tissue. A preliminary purified elastic ring from each aorta was used to determine the approximate failure load of a sample from that animal. A strain corresponding to a load that was 10% of the failure load was calculated. During forced vibration and stress relaxation testing, aortic rings from a given vessel were not extended past this strain, e10%. Forced vibration tests were performed at 1 Hz increments from 1 to 10 Hz. The amplitude of the sinusoidal extension cycles was 5% of e10%, with a maximum extension of e10%. The final 10 oscillations (of a total 100 cycles at each frequency) were used in the analysis to ensure that the samples were fully preconditioned. Force and extension data were captured at 200 Hz. For quasi-static pull-to-failure testing, elastic tissue rings were preconditioned for 100 sinusoidal extension cycles from a minimum strain of 5% of e10% to a maximum of e10%. Samples were then brought back to zero extension and extended at ~0.3 mm/s until failure. Data were recorded at 10 Hz. Analysis of Mechanical Testing Data Strain Data From Ring Samples Strain calculations corrected for the effects of the ring-shaped samples, where initial extension of the samples flattened the rings into two parallel strips before straining the elastic tissue. Therefore, the gauge length, lgauge, of the samples was taken as half of the inner circumference of the rings (representing the initial, unstrained lengths of the flattened ring’s parallel strips, measured via image analysis as described earlier). Assuming the absence of bending stiffness, strain values were thus calculated as: e¼
Dl þ li lgauge lgauge
ð3Þ
We note that Dl was the displacement of the moving actuator during testing. This was equal to the change in length of the sample once the ring was fully collapsed and extending as two parallel strips. Forced Vibration Data Force and extension were converted to stress and strain using the width and thickness measurements of each aortic elastic ring made prior to mechanical testing. Given the small mean deformations imposed
Changes in the Mechanical Properties and Residual Strain of Elastic Tissue
during the forced vibration tests, nominal stress (r) was calculated from the force data using: r¼
F F ¼ Ai 2 wi ti
ð4Þ
and strains were calculated from extension data as per Eq. (3) above. Here, F is the force (distributed across the two strips of the aortic ring as it elongates), and Ai is the total initial unloaded cross-sectional area of the two strips of the ring (calculated from its initial unloaded width wi and thickness ti). The forced vibration stresses and strains were fitted using the Levenberg-Marquardt nonlinear least squares algorithm (LabVIEW) to sinusoids of the form y = a + bÆsin(xt + c) where x is the frequency of the vibration, and a, b, and c are constants determined by the fit. The complex viscoelastic modulus (E*) was then determined from fitted sinusoids as: ro E ¼ jE jeid ¼ eid ð5Þ eo
their crown-rump length: 50–70 and 80–100 cm long, where ~100 cm crown-rump length corresponds to full term. To examine the effects of age groups, a one-way ANOVA was performed followed by Bonferronimodified t-tests for comparisons among three groups (non-pregnant and two fetal age groups). There, a significant difference was concluded when p < 0.05/3 = 0.0167. To determine the changes in any parameter during fetal development, data for each parameter were plotted as a function of fetal length and fitted with a least-squares linear regression. The regression was considered significant when p < 0.05.
RESULTS Elastic Fiber and Water Contents in Aortic Tissue From fetal to adult life, aortic elastic fiber content increased while tissue water content decreased. Expressed as percent dry weight, elastic fiber content
Here, |E*| is the magnitude of the complex viscoelastic modulus, ro and eo are the amplitudes of the fitted stress and strain sinusoids, and d is the phase shift (in radians) between them. Pull-to-Failure Data Due to the much larger deformations of the samples in the pull-to-failure tests, nominal stress was converted to true stress under an assumption of constant tissue volume during deformation: r¼
F F L FðDl þ li Þ ¼ ¼ At V Ai lgauge
ð6Þ
From the quasi-static stress–strain curve to failure, the following mechanical parameters were obtained: (i) the ultimate tensile strength (rmax) of the tissue was taken as the maximum stress achieved; (ii) the extensibility (emax) was the strain at which the tissue reached rmax; (iii) the toughness (energy absorbed per unit volume of tissue before failure) was taken as the area under the stress–strain curve up to emax; (iv) the quasi-static incremental modulus (E) was calculated as the slope of the stress–strain curve over the range 50 ± 10% of rmax. We note that this stress range is higher than that under which the dynamic viscoelastic moduli were calculated in forced vibration tests (~10% of rmax). Statistical Analysis Results are expressed as mean ± SEM and the n value used in the calculations is the number of animals providing samples for each group. Fetal data were sub-divided into two age groups estimated from
FIGURE 2. Changes in the percent composition of intact aortic tissue. (a, c) Elastic fiber content (as percent dry weight) and water content (percent wet weight) as functions of fetal crown-rump length (100 cm term),50 along with measurements from adult animals (~1–2 years old). Regression lines with 95% confidence intervals are shown for the fetal data. (b, d) Comparison of values (means 6 SE) between age groups. n 5 6, 5, and 4 for the 50–70 cm fetus, 80–100 cm fetus, and adult groups respectively. Values labeled with the same letter were not significantly different.
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rapidly increased with increasing fetal length during gestation (r2 = 0.84, Fig. 2a). There was a 21% increase between the 50–70 cm fetal length range (48.7 ± 1.0%) to the 80–100 cm fetal length range (58.5 ± 1.8%), and a further 11% increase into adulthood, reaching 64.7 ± 1.4% (Fig. 2b). The aortic water content decreased with increasing fetal length during gestation (r2 = 0.87, Fig. 2c). There was a significant drop in mean aortic water content between the 50–70 cm fetal length range (87.0 ± 0.3%) and the 80–100 cm fetal length range (83.7 ± 0.4%), and a larger drop in the transition to the adult aorta group (75.5 ± 0.4%, Fig. 2d). Aortic Dimensions The circumference of the purified aortic elastic fiber rings increased significantly during development from fetal to adult life, while the wall thinned (Fig. 3). Much of these changes were concentrated in the short, late gestational interval between the 50–70 cm fetus length range and the 80–100 cm fetus length range: approximately 60 days. Indeed, the changes here were as great as those over the much longer period encompassing postnatal development between 80–100 cm fetuses and the 1–2-year-old adult animals. While both wall thickness and circumference of the purified elastic tissue rings increased significantly with increasing fetal length (Figs. 3a and 3c), the effect on circumference was strongest (r2 = 0.85 vs. 0.45). Circumference increased by 180% between the 50–70 cm fetal length range and the 80–100 cm fetal length range, and by further 155% into adulthood (Fig. 3d). Similarly, wall thickness increased by 23% between the 50–70 cm fetal length range and the 80–100 cm fetal length range, and by further 28% into adulthood (Fig. 3b). Aortic elastic tissue ring circumference increased more than did thickness during development, resulting in a decrease in the relative wall thickness, calculated as the ratio of the thickness, h, to the unloaded radius, Ro. This ratio decreased with increasing fetal length during gestation (Fig. 3e), with mean values rapidly decreasing by 33% between the 50–70 cm fetal length range and the 80–100 cm fetal length range. There was a similar-sized decrease during postnatal development into adulthood, reaching a final value of 0.26 ± 0.02 (Fig. 3f). Residual Strain There were dramatic increases in the residual strain in the aortic elastic fiber rings throughout development, with most of this increase occurring during the
FIGURE 3. Changes in the dimensions of the purified aortic elastic tissue rings. (a, c, e) Wall thickness, inner circumference, and relative wall thickness of purified aortic elastic rings as functions of fetal crown-rump length (100 cm term),50 along with measurements from adult animals (~1–2 years old). Regression lines with 95% confidence intervals are shown for the fetal data. (b, d, f) Comparison of values (means 6 SE) between age groups. n 5 6, 4, and 5 for the 50–70 cm fetus, 80–100 cm fetus, and adult groups respectively. Values labeled with the same letter were not significantly different.
short late gestational interval. Opening angle significantly increased with fetal growth during gestation (Fig. 4a), with mean values increasing by 262% from the 50–70 cm fetal length range to the 80–100 cm fetal length range (Fig. 4b). There was, however, no significant change in residual strain from late fetal life to adulthood.
Changes in the Mechanical Properties and Residual Strain of Elastic Tissue
FIGURE 4. Changes in opening angle of purified aortic elastic tissue rings. (a) Opening angle as a function of fetal crown-rump length (100 cm term),50 along with measurements from adult animals (~1–2 years old). Regression lines with 95% confidence intervals are shown for the fetal data. (b) Comparison of opening angle values (means 6 SE) between age groups. n 5 6, 5, and 5 for the 50–70 cm fetus, 80– 100 cm fetus, and adult groups respectively. Values labeled with the same letter were not significantly different.
FIGURE 6. Changes in ultimate mechanical properties of purified aortic elastic tissue rings. (a, c) Extensibility (strain at failure) and toughness (energy per unit volume absorbed before failure) as functions of fetal crown-rump length (100 cm term),50 along with measurements from adult animals (~1–2 years old). Regression lines with 95% confidence intervals are shown for the fetal data. (b, d) Comparison of values (means 6 SE) between age groups. n 5 6, 4, and 5 for the 50–70 cm fetus, 80–100 cm fetus, and adult groups respectively. Values labeled with the same letter were not significantly different.
FIGURE 5. Representative quasi-static stress–strain curves of purified aortic elastic tissue rings from the three age groups studied: a 69 cm fetus, a 100 cm fetus (near-term), and an adult.
Mechanical Properties There were striking changes in the static and dynamic mechanical properties of the purified aortic elastic fiber rings during late gestational and postnatal development. During development, from the youngest fetal group studied to the adults, the quasi-static stress–strain curves shifted leftward—that is, equivalent stresses producing less strain. Representative curves for the three age groups are shown in Fig. 5. While the tensile strength of the elastic fiber rings remained unchanged throughout the developmental period studied (overall mean rmax = 234 ± 12 kPa, data not shown), the corresponding extensibility (strain at rmax) decreased by 50% during late gestation, but
was then unchanged (p = 0.13) into adulthood (Figs. 6a and 6b). Similar changes were observed for toughness (energy absorbed to failure, Figs. 6c and 6d). The quasi-static modulus, E, the mid-stress-range stiffness of the elastic tissue rings before failure, significantly increased during development and then from fetal to adult life. It significantly increased with increasing fetal length during gestation (Fig. 7a), while mean values increased by 66% between the two fetal length range groups (117 ± 8 and 194 ± 8 kPa respectively). This was followed by a further 72% increase over the longer postnatal interval into adulthood (333 ± 34 kPa) (Fig. 7b). As expected for viscoelastic materials, the magnitude and phase of the viscoelastic modulus increased slightly as the forced vibration frequency increased from 1 to 10 Hz (data not shown). This was observed in all three age groups. Since the forced vibration data was least variable at lower frequencies, data obtained
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FIGURE 7. Changes in stiffnesses of purified aortic rings. (a) Quasi-static modulus (E) as a function of fetal crown-rump length (100 cm term),50 along with measurements from adult animals (~1–2 years old). Regression lines with 95% confidence intervals are shown for the fetal data. (b) Comparison of values (means 6 SE) between age groups. n 5 6, 4, and 5 for the 50–70 cm fetus, 80–100 cm fetus, and adult groups respectively. Values labeled with the same letter were not significantly different.
FIGURE 8. Changes in viscoelastic behavior of purified aortic rings. (a) Viscoelastic phase angle, h, as a function of fetal crown-rump length (100 cm term),50 along with measurements from adult animals (~1–2 years old). Note that the regression line has a slope not significantly different from zero. (b) Comparison of values (means 6 SE) between age groups. n 5 6, 3, and 3 for the 50–70 cm fetus, 80–100 cm fetus, and adult groups respectively. Values labeled with the same letter were not significantly different.
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consequence of the concave-upward shape of the stress–strain curves (Fig. 5). Interestingly, the aortic elastic fiber rings became progressively less viscoelastic (that is, more elastic) with increasing age, from the youngest fetal group to the adults (Fig. 8a). The phase angle of the dynamic viscoelastic modulus (h) decreased by 40% from that for the youngest fetuses (50–70 cm fetal length range, 4.8 ± 0.4) to that for the adults (2.9 ± 0.1). As was observed with both moduli, E and |E*|, as much decrease in h was observed during late gestation as followed during the much longer postnatal interval studied (Fig. 8b).
DISCUSSION This study has demonstrated that the elastic tissue of the aorta—the structural component critical to its mechanical function as a pressure vessel and Windkessel—undergoes dramatic structural and mechanical alterations during late gestational and postnatal development. As elastic fiber content increased during development, the circumference and thickness of the purified aortic tissue both increased; however, circumference rose faster than wall thickness, causing a relative thinning of the elastic tissue. With these dramatic structural changes in aortic elastic tissue came marked alterations in its intrinsic mechanical properties. Its stiffness increased while its viscoelastic character decreased during late gestational and postnatal development. Much of these structural and mechanical changes were concentrated during late gestational development. The magnitudes of the changes over about 60 days late gestation were often as great as the changes that occurred during the much longer postnatal period (1–2 years in this study). All this evidence supports very rapid remodeling of elastic tissue occurs during late gestation. Compositional Changes in Aortic Elastic Tissue During Development
at the lowest frequency (1 Hz) was used for statistical comparisons. In comparisons between age groups, the magnitude of the complex viscoelastic modulus, |E*|, showed the same trend as was observed for the quasistatic modulus: it was lowest in the 50–70 cm fetal group and increased with increasing fetal length during gestation (data not shown). Mean values increased significantly between the two fetal length range groups (18 ± 2 to 43 ± 7 kPa) and again in the adults (80 ± 5 kPa). As expected, these stiffnesses are much lower than those calculated under much higher stresses as the quasi-static modulus, E. This is a direct
The present results showing aortic elastic fiber content increasing with development agree with previous studies.6,57 It is interesting that these changes occurred mainly during late gestation such that the elastic fiber contents of the term fetuses (100 cm fetuses) were close to those in the adults (Fig. 1), suggesting rapid structural remodeling including the accumulation of amorphous elastin during late gestation. As well, decreases in water content during fetal and postnatal periods have previously been observed in the bovine nuchal ligament9,47 and the ovine aorta.57 While the hydration and proportion of elastic tissue in
Changes in the Mechanical Properties and Residual Strain of Elastic Tissue
the aorta both dramatically changed during development, the changes did not occur on the same timelines and are not likely related: e.g., the increase in elastic fiber content occurred mainly during fetal development, whereas the water content showed its largest decrease postnatally. Changes in the Mechanical Properties of Aortic Elastic Tissue During Development Where comparisons are possible, the present mechanical results agreed well with previous studies. Quasi-static modulus values for aorta elastic tissue ranged from 117 ± 8 kPa for the younger fetal age group, to 334 ± 34 kPa for the adult elastic tissue in this study. These values were comparable in magnitude—and showed similar changes with age—to the low stress modulus values for whole aortic tissue (the low stress modulus is dominated by elastic fibers) obtained by Wells et al.57 for the developing ovine aorta. Their values ranged from 82 ± 10 kPa for fetal lambs to 144 ± 10 kPa for adults. Opening angle results for the adult aortic elastic tissue (h = 19.7 ± 4.0) were also similar to the values obtained by Gratzer and Lee22 for young, mature aortic elastic tissue rings from cattle (h = 26.5 ± 1.2). Not only did the aortic content of elastic tissue increase during development, but results from this study also demonstrated that the stiffness of that elastic tissue increases. Thus, the increasing low-stress wall stiffness observed by Wells et al.57 in the developing intact aorta, was likely due not only to an increase in elastin content but also to an increase in the intrinsic mechanical stiffness of elastic fibers (expressible in material properties like moduli). Changes in the intrinsic mechanical properties of elastic fibers themselves may not be surprising given their extensive transformation from a microfibrillar scaffold to an elastin–microfibril composite,15,40,42 along with potential increases in elastin crosslinking.3,54 We have also demonstrated a striking, developmental decrease in the extensibility of aortic elastic tissue from fetal to adult life, with much of this decrease occurring during late gestation. Aortic elastic tissue stress–strain curves shifted markedly to the left over a relatively short (~60 days) interval, with strains at fracture (extensibility) decreasing by more than a factor of 2 (Figs. 5 and 6b). This is quite opposite to the trend seen in whole aortic tissue by Wells et al.,57 where circumferential aortic extensibility in sheep increased during development from fetal to adult life—there with the stress–strain curves shifting to the right. The discrepancy between these observations—less extensible elastic lamellae within an overall more extensible aortic tissue—must therefore
be explained by remodeling of the aortic collagen since the composite mechanics of the aortic wall is largely derived from the combined stiffness of the elastic tissue and collagen components. The previous study by Wells et al.57 associated a decrease in aortic tissue viscosity from fetal to adult life with a concurrent decrease in aortic smooth muscle cell content—supporting the long-held theory that arterial cells are the primary contributors to arterial viscous losses. In this study, however, the cells were removed by the autoclaving treatment. Thus, the developmental decrease in viscosity of the purified elastic tissue cannot be attributed to changing cell content. We note that only whole tissue water content was measured, elastic fiber water content (i.e., the water content after autoclaving) was not measured. Although the water content of the elastic fibers would not be the same as the whole tissue elastic content (some water would be associated with collagen and with water-soluble proteoglycans which would be disrupted by autoclaving), the changes in elastic fiber water content should follow the same temporal pattern of change as the whole tissue water content. Wall Stresses and Remodeling of Aortic Elastic Tissue The growth and remodeling of aortic elastic tissue is probably driven by the developmental increase in physiological wall stress during fetal and postnatal development: changes associated via the Law of Laplace with the gradual rise in blood pressure and increase in vessel dimensions.28,55,57 A rise in physiological operating stress would be expected in the present bovine model: aortic circumference was increased by 180% during late gestation (Fig. 3d), and bovine fetal blood pressure increases by approximately 65% over the same interval.46 Further increases in blood pressure and aortic diameter postpartum would continue the increase of aortic wall operating stress into adulthood. An increase in aortic elastic fiber content during the developmental increase in aortic wall stress is in agreement with previous studies that have correlated increased medial wall stress during development with elastin accumulation.7,32 Indeed, Looker and Berry34 demonstrated that the bulk of elastin synthesis is directed toward the formation of new elastic lamellae, and the number of elastic lamellae is proportional to the medial wall stress. Herein, we have demonstrated a thickening of the isolated aortic elastic tissue during development, reflecting the addition of elastic lamellae—tissue deposition in the radial direction—to support increases in wall stress. In addition to the increase in aortic elastic fiber content and wall stress during development, we
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observed a dramatic increase in the residual strain exhibited by this tissue, mainly during late gestation. In a review, Gleason et al.20 emphasized ‘‘the need to discover how the deposition of elastin generates residual stress’’ in developing arteries. This is the first study, to our knowledge, to demonstrate changes in the residual stress borne by the elastic tissues of the developing fetal aorta. Similar increases in whole tissue aortic residual stress (seen as increased opening angle) has been observed during postnatal development in the rat.23 These studies and ours suggest that whole aortic tissue residual stresses may be held mainly by its elastic lamellae and are thus be modulated through elastic tissue remodeling. The ‘‘stress-growth law’’ emerging in soft tissue mechanics predicts that while global arterial growth can alter the magnitude of the average wall stress, local differential growth and/or resorption of tissue may alter the distribution of stress across the wall.16,44 This differential remodeling—as with global arterial remodeling27—is triggered by changes in mechanical loading conditions.17,18 In the present study, we have observed rapid increases in the opening angle (increased residual strain and stress) of purified aortic elastic tissue. This suggests that, during aortic development, the growth rate of elastic lamellae on the intimal side of the vessel wall may exceed that on the adventitial side—a remodeling mechanism similar to that seen in hypertension.18,39 This hypothesis is supported by histological studies on developing aortas that show more highly developed and thicker elastic lamellae in the inner layers of the aortic wall,14,41 suggesting that the inner lamellae form more rapidly or earlier during gestation. It is important to note, however, that (i) increasing blood pressure during development elicits different responses than does (ii) hypertension in mature arteries. While both developing and mature vessels thicken in order to normalize the circumferential wall stress, developing arteries do so by adding additional elastic lamellae and interlamellar units while mature arteries thicken by addition of tissue in existing interlamellar units,58,59 with little or no elastin deposition.24 Arteries are thought to develop or modulate residual stresses in order to minimize the stress gradients across the wall under physiological conditions,8,16 thereby creating a more homogeneous mechanical environment for smooth muscle cells. Increased residual strain (opening angle) over the period of development in the current study suggests that a transmural stress gradient persists in the elastic lamellae, with higher stress at the inner wall driving more rapid elastic tissue growth at this location. This hypothesis is supported by Rachev et al.,45 who
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calculated higher stresses at the inner aortic wall during development, which they point to as the trigger for enhanced growth at this location. Presumably, as the maximum adult aortic wall stress is achieved, the aortic opening angle will cease to increase as the transmural stress gradient is minimized or reduced, and homeostasis is achieved.
SUMMARY AND CONCLUSIONS This study has simultaneously shown links between development, alterations in pre-stress, and the resulting changes in mechanical properties of elastic tissue during critical late gestational development when elastic fibers are formed. This remodeling is crucial to the lifetime mechanical performance of elastic fibers and is very likely part of the signaling mechanism that makes this tissue an essential determinant of arterial morphogenesis in fetal development. The dramatic changes observed here over the last trimester of bovine gestation highlight the fact that, while both support blood pressure and perform the Windkessel function, the developing and mature aortic elastic tissues are quite different materials. While the current study examined bovine fetuses in their last trimester of gestation, this difference is certain to be magnified earlier in gestation where the elastic fibers are exclusively microfibrillar.15 Future studies will focus on associated developmental changes in elastic fiber microstructure and will attempt to elucidate the contribution of microfibrillar network to the developing elastic fiber. We hypothesize that the developmental increases in aortic physiological operating stress is a major factor driving this remodeling process. The growth and remodeling of aortic elastic tissue during this important stage in development likely serves many roles: (i) reinforcing the wall as it gradually holds higher physiological wall stress, (ii) modulation of residual stress for more uniform stress state across the aortic wall, and (iii) modulating the Windkessel function of the aorta for increases in cardiac output after birth.
ACKNOWLEDGMENTS The authors wish to thank Dr. J. Michael Lee for his valuable input on the manuscript, and O.H. Armstrong Food Services, Inc., for the donation of bovine tissues. Operational (SMW) and scholarship (EJW) funding was provided by the National Sciences and Engineering Research Council of Canada (NSERC).
Changes in the Mechanical Properties and Residual Strain of Elastic Tissue
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