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Am J Physiol Heart Circ Physiol 292: H2754-H2763, 2007. First published January 19, 2007; doi:10.1152/ajpheart.01108.2006
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Effect of elastin degradation on carotid wall mechanics as assessed by a constituent-based biomechanical model

E. Fonck,1 G. Prod'hom,1 S. Roy,1 L. Augsburger,1 D. A. Rüfenacht,2 and N. Stergiopulos1

1Laboratory of Hemodynamics and Cardiovascular Technology, School of Life Sciences, Swiss Federal Institute of Technology, Lausanne, Switzerland; and 2Neuroradiology, Radiology Department, Geneva University Hospital, Geneva, Switzerland

Submitted 11 October 2006 ; accepted in final form 17 January 2007


   ABSTRACT
TOP
 ABSTRACT
METHODS
 RESULTS
 DISCUSSION
 GRANTS
 REFERENCES
 
Arteries display a nonlinear anisotropic behavior dictated by the elastic properties and structural arrangement of its main constituents, elastin, collagen, and vascular smooth muscle. Elastin provides for structural integrity and for the compliance of the vessel at low pressure, whereas collagen gives the tensile resistance required at high pressures. Based on the model of Zulliger et al. (Zulliger MA, Rachev A, Stergiopulos N. Am J Physiol Heart Circ Physiol 287: H1335–H1343, 2004), which considers the contributions of elastin, collagen, and vascular smooth muscle cells (VSM) in an explicit form, we assessed the effects of enzymatic degradation of elastin on biomechanical properties of rabbit carotids. Pressure-diameter curves were obtained for controls and after elastin degradation, from which elastic and structural properties were derived. Data were fitted into the model of Zulliger et al. to assess elastic constants of elastin and collagen as well as the characteristics of the collagen engagement profile. The arterial segments were also prepared for histology to visualize and quantify elastin and collagen. Elastase treatment leads to a diameter enlargement, suggesting the existence of significant compressive prestresses within the wall. The elastic modulus was more ductile in treated arteries at low circumferential stretches and significantly greater at elevated circumferential stretches. Abrupt collagen fiber recruitment in elastase-treated arteries leads to a much stiffer vessel at high extensions. This change in collagen engagement properties results from structural alterations provoked by the degradation of elastin, suggesting a clear interaction between elastin and collagen, often neglected in previous constituent-based models of the arterial wall.

elastin; elasticity and stiffness; structural properties; constitutive relations; strain energy function

ELASTIN IS A PROTEIN THAT coils and recoils like a spring within the elastic fibers of connective tissue and which accounts for the elasticity of structures such as the skin, blood vessels, heart, lungs, intestines, tendons, and ligaments (25). Elastin functions in connective tissue together with collagen providing elasticity, and, alternatively, collagen provides rigidity. Elastin degradation rate strongly depends on age (38), amount of elastase present, hormones, stress, smoking, and ultraviolet exposure (39).

Elastase-treated arteries have offered an attractive model to better access the biomechanical properties of arteries after elastin degradation, as seen in major artery pathologies, such as aortic stiffening with age and aneurysm formation (8–10, 19, 26). Past studies have described elastolysis applied to ex vivo human arteries to develop mathematical models describing stress distribution postdegradation of elastin in arteries (33, 38). The biomechanical effects of elastin degradation on the extracellular matrix (ECM) and the biomechanical properties of the wall have not been discussed in detail. This is mainly due to the fact that most of the earlier biomechanical models of the arterial wall assumed uniform and homogeneous properties across the wall, thereby not allowing for the assessment of the individual contribution of each structural element of the wall [collagen, elastin, and vascular smooth muscle cells (VSM)]. Lately, significant effort has been put in developing constituent-based strain energy functions (SEF) for describing the properties as well as the three-dimensional strain and stress field in the arterial wall (13, 21, 42, 44). These SEFs account for the individual contribution of the wall's main structural elements (elastin, collagen, and VSM), providing thus a better basis for the analysis of arterial wall mechanics in health and disease. Experimental data from arteries, where selectively one or more of the structural wall elements is altered, are very useful in deriving the appropriate forms of these SEF and understanding the functional coupling between these structural elements.

This report focuses on the biomechanical properties of the arterial wall before and after enzymatic degradation of its elastin component. In specific, the goal of the study is to characterize the effect of degraded and nonfunctional elastin on vessel geometry, structural properties, elastic properties, and collagen engagement. The biomechanical analysis accomplished will help in understanding the role of elastin on the passive structural properties and in accessing its effects on collagen engagement, which is the main element kept after full elastin degradation.


   METHODS
TOP
 ABSTRACT
 METHODS
RESULTS
 DISCUSSION
 GRANTS
 REFERENCES
 
Biological Material

Arterial specimens. Left common carotid arteries were harvested from 12 White New Zealand female rabbits (3.0 ± 0.2 kg) from a local slaughterhouse. Before and after the arteries were excised, longitudinal length was measured to assess the physiological in situ longitudinal stretch ratio. The arteries were transported in PBS with anti-inflammatory reagents to the laboratory at 4°C. Upon arrival, arteries were cleaned from extraneous tissues with great care. The common carotid artery was typically 50–60 mm in length, out of which two segments of ~16 mm in length were cut. One was used as a control artery and the other for elastase treatment, with the order, proximal or distal, being randomly assessed before each experiment.

Biochemical Treatment

Enzymatic degradation. Elastase type, concentration, incubation time, and its inhibitor were determined experimentally, using as basis the published works of Miskolczi et al. (31) and Greenwald et al. (18). We used high-purity elastase enzyme (6,850 U/mg; Elastin). Elastase was added into the bath, where laid the physiologically stretched artery, to obtain a final concentration of 6.85 U/ml in solution. The artery was left 30 min at 37°C in the solution under continuous manual stirring and thereafter washed with fresh PBS for a few minutes. The artery bath was then replaced with warm (37°C) PBS containing 2 U/ml aprotinin (Elastin) to inhibit the elastase trapped in the arterial wall. The bath was emptied 30 min later and filled up with warm PBS preceding immediate wash of the artery with PBS. Earlier experiments (42, 44) and preliminary tests on rabbit carotids have verified the viability and integrity of VSM left in PBS for the entire duration of the control experiments. Note that arteries were left mounted on the cannulas throughout the entire enzymatic procedure, avoiding hazardous manipulations.

Biomechanical Analysis

Geometrical analysis. Before the arteries were mounted on the apparatus for biomechanical testing, a small ring (~3 mm) was cut off from each arterial segment for measurement of the opening angle, wall thickness, and inner and outer diameters. The rings were examined with a light microscope at x20 magnification (Axioplan2; Zeiss). Each ring was cut open and let to rest for 30 min in a warm PBS dish. The cut open ring was then observed with the same microscope to measure the opening angle. The measurements of the internal diameter, external diameter, arterial wall thickness, and the opening angle of the arterial ring were obtained using ImageJ 1.32 software (National Institutes of Health, Bethesda, MD).

Biomechanical testing. Arteries were mounted on an inflation/extension testing device, where a bath was filled with PBS maintained at 37°C. The arteries were stretched to their physiological stretch ({lambda}z), typically at ~1.4 times their resting elongation. Arteries were then submitted to preconditioning cycles (5 cycles on average) composed of a series of inflations and deflations ranging from 0 to 180 mmHg at an approximate rate of 1.5 mmHg/s by means of a syringe pump (model sp201iw; World Precision Instruments; see Ref. 43). Pressure-diameter inflation and deflation was repeated to ensure repeatability of the curve and to minimize viscous effects (16). External diameter data were measured with a digital high-accuracy charge-coupled device (CCD) micrometer (model LS-7030; Keyence). Inflation pressure was recorded with a blood pressure transducer (BLPR; World Precision Instruments). Pressure and external diameter data were collected on a data acquisition card (DAQ 6285; National Instruments, Austin, TX) and visualized on a data acquisition interface (7.1 Labview; National Instruments). The passive arterial state (control state) was obtained with the addition of sodium nitroprusside (SNP, 10–4 M) before each recorded measurement and after partial removal of elastin of the arteries. To verify that the applied dose of SNP totally abolishes vascular smooth muscle tone, we compared the pressure-diameter curves with those obtained after treatment with cytochalasin D, which destroys the VSM cytoskeleton, thus preventing any development of vascular smooth muscle tone. The results showed that treatment with SNP and cytochalasin D yielded almost identical pressure-diameter curves, suggesting that administration of 10–4 M of SNP leads to a totally passive wall state. The whole procedure was then repeated for a longitudinal stretch of {lambda}z = 1.6.

Arterial Wall Properties

Structural properties. External diameter-pressure data were combined with the zero-load data and processed to provide arterial wall thickness, wall area, compliance, and elastic modulus vs. circumferential stretch. For comparing diameters at defined pressure, the following empirical equation was used to fit the pressure-diameter curves

Formula 1(1)
where dout is the external diameter, p the pressure recorded at dout, and a, b, c, and d are constants. This empirical model allows for a good fit of the sigmoidal pressure-diameter curves and is superior to the often-used model of Langewouters et al. (27) in terms of quality of fit and faithful representation of slopes at both low- and high-pressure regimes. Using the wall thickness, h0, and the outer diameter, d0, measured at the zero load state, one can easily compute the wall area, Aw, at the loaded state

Formula 2(2)
Assuming incompressible arterial wall, thickness [h(p)] and internal diameter (din) are expressed as

Formula 3(3)
Area compliance (CA) was computed as

Formula 4(4)
where A(p) is the luminal area of the arterial segment.

Elastic properties. The elastic properties of the arterial wall are expressed using the Hudetz incremental elastic modulus (22) by

Formula 5(5)
where dmid is the midwall diameter (dmid = dout – h). The Hudetz modulus is an indicator of the elastic properties of the wall with limited scope, since it refers only to the apparent material properties in the circumferential direction.

Histological Analysis

Elastic fibers through light microscopy. Sections for histological analysis were taken from both nontreated and elastase-treated arterial segments. Segments were mounted on cannulas, pressurized at 100 mmHg, and dropped in a 4% formalin solution for 2 h. Segments were then dehydrated in graded alcohol and xylol, embedded in paraffin wax from which sections 5 µm thick were cut on a microtome (Leica EG 1120), and divided into two subgroups. The first subgroup was stained with Miller's Elastic stain for elastic fibers, and the second subgroup was prepared for polarized light microscopy and confocal laser microscopy using 0.05% Picro Sirius red. The sections stained with Miller's Elastic stain were examined by light microscopy (Axioplan2; Zeiss), and digital images were obtained by a commercial imaging system (AxioCam HRm and AxioVision Rel 4.2; Zeiss) at x20 magnification.

Elastin structure through confocal microscopy. Elastin fiber natural autofluorescence in the green can easily be observed with confocal laser microscopy. Confocal images were acquired on a Leica TCS SP2 AOBS microscope with an excitation wavelength of 488 nm from an Argon laser with a bandwidth set at 500–575 nm for the elastin. Collagen was visualized in the red spectra with an Argon laser at 514 nm for the excitation and a 525- to 600-nm bandwidth for the emission. Collagen autofluorescence was enhanced significantly after 0.05% Picro Sirius red staining. Sections were observed under a HCX PL APO 63x/1.4 immersion objective (Leica, Wetzlar, Germany).

Collagen structure through polarized light microscopy. Birefringence optic microscopy was performed on prepared carotid sections with an Olympus AX 70 and UPlan FL20x/0.5 objective (Olympus, Tokyo, Japan), and pictures were acquired with an Olympus DP 70 CCD color camera. As mentioned by Montes (32) and Humphrey and Canham (24), this procedure allows examination of the distinct structural types of collagen and to deduction of any collagen degradation resulting from the enzymatic treatment. Semiquantitative analysis was performed on birefringence images for collagen distribution changes. Type I collagen has been localized in blood vessels in the form of thick, brilliant, and strongly birefringent yellow or red fibers. On the other hand, type III collagen displays a weak birefringence of a greenish color (32). Histology images were initially processed using Adobe Photoshop CS (Adobe System, San Jose, CA) to delineate the region of interest. This region was then defined as a unit area, also referred to as the reference surface, to further compute the collagen relative density. An image analysis algorithm was developed on MetaMorph 6.3 (Molecular Devices) to compute the relative proportion for collagen I and III. The signal intensity for the red, green, and yellow layers was threshold at 80% of total signal strength for each three layers to reduce background artifacts. The structure on the images were then analyzed, and the results are expressed as the fraction of collagen I and III to the total signal measured on the birefringence slides, considered on polarization as the total collagen content. This total birefringent signal was also given as a fraction of the reference surface to express the collagen abundance in the wall media.

Mathematical Model

SEF. We based our analysis on the structural SEF proposed by Zulliger et al. (42), which assumes that elastin and collagen act in parallel and their contribution to the total SEF is additive. Because vascular smooth muscle is totally relaxed, we have considered only the passive properties of the arterial wall. The SEF is thus separated into an isotropic and a transversely isotropic part representing the elastin and collagen, respectively:

Formula 6(6)
Each part is weighted by the fractions of the total wall cross-sectional area occupied by elastin, felast, and collagen, fcoll, so that the SEFs Welast and Wcoll result in effective stresses for each component. In previous work (42), it was found that a good description for the SEF of elastin Welast is:

Formula 7(7)
where the constant celast > 0 is the parameter describing the elastic properties of elastin, I1 = 2(Er + E{theta} + Ez) + 3 is the first invariant of the Cauchy-Green strain tensor, and where Ei = 1/2 ({lambda}i2 – 1), i = {theta}, z, r (37).

In our study, we assume that the collagen fibers, in the media, are oriented circumferentially, a simplification based on previous observations (7, 13, 14, 17, 40). We keep in mind that the collagen fiber distribution circumferentially may be specific to human cerebral arteries, since other authors have reported higher angles for different vessels (20). We further assume that the engagement of the collagen as the artery is stretched and inflated is distributed in some statistical manner (6, 40, 41). Following Zulliger et al. (42), we have applied the log-logistic PDF (probability density function, {rho}fiber), describing the statistical distribution of the circumferential strain E{theta} at which collagen starts bearing load:

Formula 8(8)
where b > 0 is a scaling parameter, k > 0 is a parameter defining the shape of the PDF, and E0 is the circumferential strain bound, for which all collagen fibers do not bear load. Setting E0 = 0 prevents collagen fibers from being loaded and thus exerting a force when the tissue is in its zero stress state (ZSS). Zulliger et al. admit that "the choice of the collagen fiber engagement distribution being log-logistic is fairly arbitrary." However, Zulliger et al. have tried a number of different functions as possible candidates for describing the PDF, and using as a criterion the quality of the fit on a series of experimental data they found that the log-logistic function gave the best results. An individual collagen fiber is described by the SEF as:

Formula 9(9)
where ccoll is the elastic constant associated with the collagen, and E' is the local strain in the direction of the fiber. Although somewhat arbitrary, ccoll was set equal to 200 MPa in accordance to earlier publications and with the goal to minimize the number of unknown parameters. To describe the ensemble of circumferentially oriented collagen, we can convolute the individual fiber SEF, Wfiber, with the fiber engagement PDF, {rho}fiber:

Formula 10(10)
and equation 6 now becomes:

Formula 11(11)
The equations used for deriving local radial, circumferential, and longitudinal stresses and the diameter values at a given pressure are given in Zulliger et al. (42). The experimental pressure-diameter curves were fitted using the model of Zulliger et al. The fit was driven by the minimization of the sum of squares of the difference between experimental and theoretically derived diameter at each level of pressure. The parameters derived from the fit were the elastin constant, celast, and the collagen PDF parameters b and k.

Statistics

All reported values are given as means and SD for the control group (control, n = 6) and elastase-treated group (treated, n = 6). Two-tailed Student's t-tests were performed for comparisons between the means of diameter and compliance at each level of pressure, and one-way ANOVA test was conducted for the incremental elastic modulus analysis. The level of significance was taken when P < 0.05.


   RESULTS
TOP
 ABSTRACT
 METHODS
 RESULTS
DISCUSSION
 GRANTS
 REFERENCES
 
Histology

Histological sections of control and elastase-treated arteries are shown in Fig. 1, A and B. The sections stained with Miller's Elastic stain show clearly the elastin degradation in elastase-treated arteries compared with control arteries. Indeed, the histology of the control arteries (Fig. 1A) shows densely packed parallel elastin fibers arranged uniformly as opposed to disrupted and much scarcer elastin sheets of the treated arteries (Fig. 1B). The effects of elastase are very clearly seen on the internal elastic lamina, where the disrupted layer is now seen as wavy fragments covering only a fraction of the internal circumference. On Fig. 1, C and D, we can clearly distinguish the elastin laminas in bright green and collagen fibers in red from confocal microscopy. The effects of elastase result noticeably on the fragmentation of the elastin into loose elastin elements. Figure 1B seems to show that elastin is diminished, but Fig. 1D does not suggest any decrease in elastin content, but mostly elastin fragmentation, which is also seen in Fig. 1B. Because Fig. 1D results from elastin autofluorescence, we tend to believe that elastin is indeed present in substantial quantity after elastase, as seen in Fig. 1D, and that the primary effect of elastase is fragmentation. Differences seen between Figs. 1, A and B, may be, in part, attributed to differential contrast and not really to lack of (fragmented) elastin in Fig. 1B.


Figure 1
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  		Fig. 1. Histological sections of the carotid arterial wall showing elastin fibers (A and B) and collagen fibers (C and D) before (A and C) and after (B and D) enzymatic degradation. Blue-stained elastin fibers arranged in parallel to the inner surface in controls are clearly visible in A. After the elastase treatment, only fragmented elastin fibers are visible in B. Confocal microscopy differentiates elastin fiber (bright green) from collagen (red) before (C) and after (D) elastase treatment and shows the fragmented elastin (green) and the collagen fibers (red).

 
Figure 2, A and B, shows polarized light microscopy pictures, allowing visualization of the different collagen fibers of two segments taken from the same artery, one segment kept normal as the other was treated with elastase. Semiquantitative analysis revealed that, in control arteries, the relative fractions of collagen I and III were 88 and 12%, respectively. For the elastase-treated arteries, the relative fractions of collagen I and III were remarkably similar to the control arteries (Fig. 2C).


Figure 2
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  		Fig. 2. Polarized light microscopy displaying the different type of collagen fibers. Collagen fibers of type 1 are displayed in yellow and red, and collagen fibers of type III are displayed in green in control (A) and elastase-treated (B) arteries. Semiquantitative analysis of the percentage of collagen I and III for the control and elastase-treated arteries is shown in C.

 
Opening Angle

The mean opening angle was 92.17 ± 11.43 degrees for the control arteries (Fig. 3B), and the mean opening angle for the treated arteries was significantly smaller [12.33 ± 8.80 degrees (P < 0.005; n = 6); Fig. 3D].


Figure 3
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  		Fig. 3. Microscopic images of cross sections of the carotid arterial wall before (A and C) and after (B and D) elastase treatment. E: wall thickness at the zero stress state. F: opening angle at the zero stress state.

 
Biomechanical Analysis

For all measurements, raw pressure-diameter data were fitted with equation 1, yielding excellent quality of fits (r2 > 0.99 for all arteries). Mean pressure-diameter curves are shown in Fig. 4. At all pressures, the inner diameter of treated arteries was higher than for control arteries. The difference is larger at low pressures (pressure <50 mmHg), suggesting that compressive stresses are released upon destruction of elastin, leading to diameter enlargement in the elastase-treated group. Control arteries exhibited a clear inflection point at 40 mmHg, which corresponds to the maximum of compliance (Fig. 5), and, functionally, to the initiation of engagement of collagen fibers, which progressively limits diameter expansion. Treated arteries show no inflection point but a very rapid diameter increase in the very low pressure range (0–25 mmHg) indicative of the relative absence of elastin, which constitutes the main structural element resisting inflation in this low pressure range. For pressures >25 mmHg, treated arteries are essentially flat, suggesting that a significant amount of collagen fibers have already engaged, limiting further diameter extension. In control arteries, pressure-diameter curves become essentially flat only for pressures >120 mmHg. The asymptotic maximal diameters for the control and treated arteries were 3 ± 0.2 and 3.3 ± 0.3 mm, respectively.


Figure 4
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  		Fig. 4. Pressure-diameter curves showing alterations between the control group (n = 6) and elastase-treated group (n = 6).

 

Figure 5
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  		Fig. 5. Pressure-compliance curves of the control group (n = 6) and the elastase-treated group (n = 6).

 
The mean compliance curves for both control and treated arteries are shown in Fig. 5. Control arteries show the typical biphasic compliance curves with a maximal peak at 40 mmHg of pressure. At the physiological pressure of 70 mmHg, compliance was 0.045 mm2/mmHg, which corresponds to a pulse wave velocity of 5.2 m/s, a value well within the physiological range for carotid arteries. In contrast, the compliance of treated arteries decreased monotonically and fairly abruptly in the low pressure range (pressure <30 mmHg), remaining at very low levels thereafter. For the sake of comparison, the compliance at 70 mmHg was 3.33–4 mm2/mmHg and the corresponding wave velocity 68.9 m/s, which is indicative of very stiff arteries attained only in severe forms of arteriosclerosis.

The incremental elastic modulus plotted against circumferential stretch, {lambda}{Theta}, is shown in Fig. 6. To compare the elastic properties of the two arterial groups, we used as reference dimension the zero load state of control arteries. For low levels of circumferential stretch, the wall material of the treated arteries was significantly more ductile than the control arterial wall material, which is consistent with the notion that at low stretch ratios the main wall element offering resistance to tension is elastin. At low stretch ratios ({lambda}{Theta} < 1.5), the incremental elastic modulus of the control group is progressively becoming greater, and the increase in stiffness is accelerated for stretch ratios >1.5, apparently by the higher rates of recruitment of collagen. In contrast, the elastic modulus of the treated arteries stays lower (softer material because of lack of elastin) and fairly constant until reaching a stretch ratio of ~1.85. At higher circumferential stretch (>1.85), the treated arteries quickly became stiffer in a fairly abrupt manner, pointing out an accelerated rate of engagement of the collagen fibers.


Figure 6
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  		Fig. 6. Hudetz incremental elastic modulus vs. midwall circumferential stretch ratios of the control (n = 6) and the elastase-treated group (n = 6).

 
Constituent-Based Structural Analysis

Figure 7 shows the pressure diameter of the control and treated artery groups as well as the best fit of the theoretical model evaluated from the constituent-based SEF (Eq. 6). The elastic constant of collagen, ccoll, was set to 200 MPa, a value well within the physiological range of values reported in the literature (3, 30). The elastic constant of elastin, celast, was let free for the fit, as were also the parameters of the probability density function describing the engagement of collagen, b and k. The quality of the fit was excellent for both artery groups (r2 > 0.99 and r2 > 0.95 for control and treated arterial groups, respectively). The parameters of the fit are given in the tables at the bottom of Fig. 7. The model predicts that the effective elastic modulus of elastin decreases by >90% (celast = 23,450 Pa in control to celast = 2,150 Pa in treated arteries), which, assuming that the real elastic constant of the material did not change, would mean that the effective elastin fraction in the wall decreased by ~90% after elastase treatment. Furthermore, the parameters of the collagen engagement function k and b yielding the best fit are different between control and treated arteries, suggesting that destruction of elastin affects collagen engagement characteristics. Elastase treatment led to an increase in k from 15.3 in control to 37.3 in treated arteries and a decrease in b from 0.920 in control to 0.872 in the treated artery group. The collagen engagement probability density functions for the control and elastase-treated groups are displayed in Fig. 8. Destruction and fragmentation of elastin leads to a much narrower probability density function for collagen, meaning that in treated arteries few collagen fibers seem to engage at circumferential stretch ratios <1.75, but at higher strains engagement takes place much more abruptly. To verify that changes in the collagen engagement parameters k and b are essential in predicting the pressure-diameter curves after elastase treatment, we plotted the theoretical pressure-diameter curve using the elastin constant of the treated arteries (celast = 2,150 Pa) but keeping the collagen engagement parameters k and b the same as in control (k = 15.3 and b = 0.920; Fig. 7). We observe that, although the lower elastin constant does allow for a significant diameter increase, the predicted diameter stays lower than the measured diameter values for the entire experimental pressure range, suggesting that the changes in the pattern of collagen engagement are indeed necessary for predicting the arterial wall behavior after elastase treatment.


Figure 7
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  		Fig. 7. Experimental internal radius-pressure curves for control (circles) and treated (squares) groups fitted with the theoretical model of Zulliger et al. (solid lines). The table on bottom shows the parameters derived from the best fit. The dashed line gives the theoretical curve for the treated artery group as predicted by the model of Zulliger et al. when the parameter values for collagen engagement k and b are kept the same as in control arteries (k = 15.3 and b = 0.92).

 

Figure 8
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  		Fig. 8. Probability density function for describing the engagement of collagen fibers as a function of circumferential stretch for control arteries (solid line) and elastase-treated arteries (dashed line).

 

   DISCUSSION
TOP
 ABSTRACT
 METHODS
 RESULTS
 DISCUSSION
GRANTS
 REFERENCES
 
We have treated rabbit carotid arteries with elastase to examine the effects of elastin fragmentation and digestion on vessel geometry, structural properties, elastic material properties, and collagen engagement. We have used classical biomechanical analysis to assess the wall's structural and elastic properties (pressure-diameter, compliance-pressure, and elastic modulus-stretch ratio curves). We have further analyzed the data using a structure-based SEF, which allowed us, in an indirect fashion, to assess the effective elastin destruction rate and the collagen engagement properties. We have found that elastase treatment leads to a diameter enlargement with an almost complete abolishment of the structural integrity of the artery at the low-strain range. We have also found that collagen engagement properties have been significantly affected, suggesting a clear interaction between the two main structural constituents of the arterial wall.

Effects of Elastase on Structural Properties

The enzymatic degradation of elastin affected significantly the geometrical characteristics of the artery, and this is manifested both in the pressure-diameter curves but also in the zero stress state in concordance with similar findings reported by Greenwald et al. (18). The fragmentation of the elastin, clearly evidenced in histology, led to a substantial increase in inner diameter at the low to medium pressure range. Similar observations were reported previously in earlier studies for dog carotid arteries and even for human iliac arteries (5). The diameter increase after destruction of the elastin network means that some residual compressive stresses exist, which are released when elastin is fragmented. This change in residual stresses is also manifested in the change in the opening angle (Fig. 3), although changes in opening angle are sensitive not only to the level of prestresses but also to their relative distribution across the wall thickness. Earlier works by Dobrin and Gley (1) as well as by Roach and Burton (34) have clearly demonstrated the role of elastin at maintaining the structural aspect of the arterial wall. Roach and Burton demonstrated that, at low pressures, the resistance to stretch was the result of elastin alone, whereas at high pressures it was because of collagen alone. As the collagen engages, the pressure load is redistributed, and collagen limits the arterial wall extension. Our study reconfirms these results and puts the emphasis on the significant changes in the levels of prestress caused by the elastin degradation. It is noteworthy that elastase treatment also led to an increase in axial length by ~10% (data not shown), pointing out the existence of longitudinal residual strain also released upon elastin fragmentation (12). It is also worth mentioning that elastase treatment led a slight decrease, in the order of 6%, in the wall cross-sectional area after elastase treatment, since others have reported a swelling (18).

Effects of Elastase on Material Properties

The incremental elastic modulus of the control group is greater than the incremental elastic modulus of the treated arteries over the low to medium range of the circumferential strains. This results from the fact that, at low to medium strains, collagen is not yet sufficiently engaged, and thus elastin is the main determinant of the wall elastic modulus. The decrease in the effective incremental elastic modulus after elastase treatment at low strains, however, should not be attributed to any intrinsic change in the elastic properties of elastin but rather to the decrease in the absolute amount of elastin carrying load effectively. As a large part of elastin is dissolved and fragmented and the stresses carried by the remaining elastin are still calculated by reference to the entire wall thickness, and not to the proportion of wall thickness occupied by the remaining functional elastin, the estimated value of elastic modulus drops. Alternatively, one can use the values of incremental elastic modulus after treatment to estimate, indirectly, the fraction of remaining functional elastin after treatment by assuming that the material properties of elastin have not changed. The assumption of no effect of elastase on the elastic constants of remaining elastin fibers has, however, little support from experimental evidence, and, in fact, low levels of elastase may well change the intrinsic properties of elastin by breaking down cross links.

At high circumferential stretches, the incremental elastic modulus of elastase-treated arteries increases quite abruptly for {lambda}{theta} > 1.85, whereas, in the control artery group, the elastic modulus increases much more gradually. It points out the different rates of engagement of collagen between the two groups. This difference is best exemplified on the graphs showing the statistical description of collagen engagement as a function of circumferential strain (Fig. 8). Indeed, control arteries exhibit a much wider distribution of the probability density function, meaning that collagen fibers are engaged much more gradually, and this is reflected on the gradual and smooth increase in elastic modulus over the range of 1.45 < {lambda}{theta} < 1.65. In contrast, treated arteries with a very narrow distribution of the probability density function exhibit a more rapid recruitment of collagen in the vicinity of {lambda}{theta} = 1.85, which is also reflected on the abrupt increase in elastic modulus around the same level of circumferential stretches (Fig. 6).

The above findings point out that elastase treatment may have led to changes in collagen fiber integrity. Although the elastase enzyme is of high purity, we could detect some gelatin activity in vitro (data not shown). We have found that the relative content of collagen type I and III did not change after elastase treatment (Fig. 2C), supporting the assumption that elastase is specific to elastin and does not affect differentially collagen I or III. Furthermore, the analysis of the collagen through polarized microscopy does not seem to suggest any visual impairment of collagen, at least at the light microscopic level. Hence, the observed effects on collagen engagement properties are potentially the result of a structural (orientation angle, waviness, cross-linking) rearrangement of the collagen. It is reasonable thus to assume a possible direct interaction between elastin and collagen, allowing an efficient transition of the load from elastin to collagen fibers and inducing a smoother recruitment of collagen. This interaction implies that the two main populations of scleroproteins in the arterial wall are functionally linked, especially in the transition from low (elastin-dominated) to high (collagen-dominated) strain region. We may thus conclude that earlier structural biomechanical models of the arterial wall, including the present one, which assume fully parallel and independent contributions of elastin and collagen in the biomechanical properties of the wall, present serious shortcomings and need thus to be revisited (36, 44).

Physiopathological Significance

A lack of elastin synthesis during the embryonic stage or an increase of elastin degradation is responsible for various pathologies, including premature skin aging, supravalvular aortic stenosis with or without William's syndrome, or aneurysmal formation (26, 29, 35). Elastolysis occurs in ageing and is accountable for the loss of elasticity of organs, which in turn do not function with the same efficiency (i.e., aortic compliance or skin wrinkling; see Ref. 15). For arterial tissue, in particular, loss or damage of elastin provokes severe and irreversible mechanical and biological changes within the ECM of vessels. The process is responsible for the stiffening of the arteries and the tortuous geometrical changes with ageing, as put forward by Dobrin et al. (12). Avolio et al. (2) have put forward a hypothesis that the number of repeated pulsations to which the arterial wall is subjected (age times heart rate) leads to progressive elastin damage. This loss in functional elasticity increases luminal wall stress, cardiac afterload, and endothelial damage and predisposes the vascular tissue to atherosclerosis. Furthermore, Martyn and Greenwald (29) have suggested that impaired synthesis of elastin in the aorta could lead to permanent changes in the mechanical properties and predispose an individual to higher blood pressure, increased left-ventricular mass, and cardiovascular disease. Loss of aortic elasticity with ageing is not only related to damage of elastin but also to changes in collagen structure. Many reports have elucidated the role of advanced glycation end products in augmenting arterial stiffness through collagen cross-linking (1). One of the interesting findings of our study is that elastin degradation per se alters the collagen engagement properties, thereby contributing to functional stiffening of the wall without necessitating the presence of any cross-linking agents.

Degradation of elastin may also be an initiating event in the formation and development of aortic and cerebral aneurysms (11, 23, 24). Histological examination of cerebral aneurysms shows clear fragmentation of the internal elastic lamina and almost total absence of elastin in the aneurysmal sac. It is likely that elastolytic processes by vascular cells or infiltrated inflammatory cells, as mediated by flow effects on the endothelial layer of the cerebral vessels, are a precursor to the development of aneurysms (28).

Limitations

Semiquantifications of elastin and collagen were done through histology, which permits only qualitative measurements of elastin and collagen content. Measuring opening angles and wall thickness of treated arteries was delicate due to the fact that the rings were too soft and not holding a sturdy structure, which may have introduced errors on calculating the exact zero stress state. Furthermore, we have assumed the parameter Eo to be zero for both control and treated arteries. This parameter defines the level of total strain below which no collagen fiber is engaged. It is likely that Eo = 0 is an oversimplification of reality and that elastase treatment may alter this parameter as a result of the induced structural changes within the wall. A change in Eo may impact on the observed changes in the other two parameters defining the PDF, b and k.

In conclusion, enzymatic degradation of elastin in common carotid arteries modified their structural and elastic properties. Removing elastin resulted in inner diameter enlargement and substantial release of tensile prestresses. The elastic modulus was lower in treated arteries at low circumferential stretches and significantly higher at elevated circumferential stretch levels. Abrupt collagen fiber recruitment leads to a much stiffer vessel at high extensions. Destruction of elastin induces profound changes in collagen recruitment properties, pointing out a strong functional interaction between elastin and collagen. Structural changes induced by elastin damage are of major importance and need to be understood thoroughly, since elastin degradation is present in many important physiopathological processes, such as aortic stiffening and aneurysm development.


   GRANTS
TOP
 ABSTRACT
 METHODS
 RESULTS
 DISCUSSION
 GRANTS
REFERENCES
 
This work is supported by the Swiss National Science Foundation Grant 510691 and the Swiss Federal Institute of Technology Lausanne.


   ACKNOWLEDGMENTS
 
We thank Dr. Paolo Silacci and Dr. Makoto Ohta for valuable discussions while preparing this study. We thank Michel Bachmann for technical expertise.


   FOOTNOTES
 

Address for reprint requests and other correspondence: E. Fonck, Laboratory of Hemodynamics and Cardiovascular Technology, AI 1230, Station 15, 1015 Lausanne, Switzerland (e-mail: edouard.fonck{at}epfl.ch)

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
 METHODS
 RESULTS
 DISCUSSION
 GRANTS
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