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The effect of proteolytic treatment on plastic deformation of porcine aortic tissue
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Research paper
The effect of proteolytic treatment on plastic deformation of porcine aortic tissue Jarin A. Kratzberg, Patricia J. Walker, Elizabeth Rikkers, Madhavan L. Raghavan ∗ Department of Biomedical Engineering, University of Iowa, Iowa City, IA, United States
A R T I C L E
I N F O
A B S T R A C T
Article history:
The objective of this work was to assess whether selective proteolysis of elastin and/or col-
Received 25 May 2007
lagen in a porcine aorta followed by mechanical creep loading would result in an aneurysm-
Received in revised form
like permanent tissue stretch. The underlying motivations were to (1) test the feasibility of
2 April 2008
developing an in vitro abdominal aortic aneurysm (AAA) model, and (2) understand what
Accepted 4 April 2008
role, if any, that passive creep-induced stretching plays in aneurysmal dilation. Multiple
Published online 23 April 2008
circumferentially oriented flat specimen strips were cut from the porcine thoracic aorta of ten adult pigs. Specimens were subjected to one of six treatment protocols: Untreated
Keywords:
controls (UC; N = 23), complete elastin degradation (E; N = 10), partial elastin degrada-
Aneurysm
tion (Ep ; N = 10), partial collagen degradation (Cp ; N = 22), and partial degradation of both
AAA
elastin and collagen (Ep + Cp ; N = 3). All specimens were then subjected to cyclic creep (10
Elastin
min/cycle) with increasing load amplitude until failure. The zero-load strain prior to the
Collagen
creep cycle where failure occurred was defined as load-induced plastic strain. The plastic
Enzyme degradation
strain induced by treatment alone, creep loading alone and the total was determined for all specimens. The total plastic strain was significantly greater for E (mean ± SD = 48.2 ±17.6, p < 0.005), Ep (41.6 ± 11.1, p < 0.0005), but not for Ep + Cp (48.9 ± 21.6, p = 0.17) and Cp (22.2 ± 12.8, p = 0.14) compared to UC (17.7 ± 6.1). Of the total plastic strain, treatmentinduced plastic strain was high for those specimens subjected to partial or total elastin degradation (E, Ep , Ep + Cp ), but not for those where elastin was intact (Cp ). However, loadinduced plastic strain in the treated specimens was not different in any of the treated groups compared to controls. Maximum total plastic strain of 78.6% was induced in one porcine aortic tissue from the E group. Even this is far lower than what would be needed for creating a realistic in vitro AAA model. Our findings do not support the feasibility of developing an in vitro AAA by enzymolysis followed by passive stretching of the aorta. The findings also suggest that AAA formation is unlikely to be a passive creep-induced stretching of a proteolytically degraded aortic wall as conventional thinking may suggest, but rather may be predominantly due to growth and remodeling of the aortic wall. c 2008 Elsevier Ltd. All rights reserved.
1.
Background
Abdominal aortic aneurysm (AAA) is a localized dilation of the abdominal aorta. AAA is characterized by a thinning
media with decreased elastin, thickening of the adventitia (outermost layer of the aorta) and neointima (Baxter and Halloran, 1994), and an increase in proteolytic enzyme (collaganase and elastase) synthesis. AAA pathogenesis is
∗ Corresponding address: Department of Biomedical Engineering, 1422 Seamans Center (SC), University of Iowa, Iowa City, IA 52242-1527, United States. Tel.: +1 319 335 5704; fax: +1 319 335 5631. E-mail address: [email protected] (M.L. Raghavan). c 2008 Elsevier Ltd. All rights reserved. 1751-6161/$ - see front matter doi:10.1016/j.jmbbm.2008.04.001
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thought to be triggered by enzymatic degradation of elastin combined with repetitive pressurization of the aorta and remodeling of collagen (Baxter et al., 1994; Dobrin, 1983; Keen and Dobrin, 2000; MacSweeney et al., 1994). A reliable animal model of AAA can vastly aid research toward understanding this disease. Researchers have been successful in developing in vivo aneurysm models in small animals such as rats, mice, and rabbits (Andrews et al., 1975; Anidjar et al., 1990; Gertz et al., 1988; Yang et al., 2007) by perfusing the abdominal aorta with elastase – a proteolytic enzyme that can selectively digest elastin. Such small animal models have been valuable in studying genetic and histological aspects of AAA, but not biomechanical characteristics or graft implant function owing to their miniature size. There has been only one report of similar success in large animals. Boudghene et al. (1993) reported successfully inducing a small dilation in dog abdominal aorta in vivo using the elastase perfusion method for the purpose of testing endovascular grafts. Similar attempts in other animals have not been successful (Marinov et al., 1997). The ability to create reliable large animal models of AAA remains elusive. Dadgar et al. (1997) proposed the development of an in vitro AAA model using harvested porcine thoracic or abdominal aorta. In this preliminary study, they showed that the excised porcine aorta is weakened by exposure to collagenase – proteolytic enzyme that selectively digests collagen – and that the level of this weakening may be controlled by exposure time. They hypothesized that a weakened aorta may then be biomechanically stretched in a laboratory to induce an aneurysm-like dilation, thus creating a dilated aorta with the morphological characteristics of a human AAA. Earlier, in an in vitro study, Dobrin and Mrkvicka (1994) reported that harvested tubular human iliac artery dilated in vitro when exposed to elastase, but ruptured when exposed to collagenase under physiologic pressures. Miskolczi et al. (1997) reported that porcine carotid artery dilated slightly at its bifurcation when exposed to elastase and subjected to pulsatile pressure in a saline bath. Neither of these studies (Dobrin and Mrkvicka, 1994; Miskolczi et al., 1997) reported the level of zero-load dilation achieved, however small they were. More recently Davis et al. (2005) reported plastic deformation in porcine common carotid arteries due to axial loading. This is an important discovery, as it shows that passive loading of an artery can cause permanent deformation. However, the feasibility of developing an in vitro AAA model through the combination of passive loading and enzymatic treatment of healthy aortic tissue has not been addressed. An in vitro AAA model would be a valuable tool for studying the effects of dilation on tissue properties, but it is unlikely to be of value in studying histological or genetic aspects of this disease since it does not accommodate the active remodeling that is undoubtedly taking place in vivo. Perhaps the most valuable utility of an in vitro AAA model is in testing new designs of endovascular grafts. Graft delivery designs, graft-to-aorta attachment designs and stent architecture designs may all be studied in highly controlled environments with in vitro AAA models. Further, fatigue testing of an endovascular graft implanted in an in vitro AAA would be substantially more realistic than current
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approaches where compliant cylindrical tubes are used. An in vitro AAA model can also be valuable in predicting shortterm implant performance such as post-procedural endoleak, although it is unlikely to be a good model for long-term complications. Secondly, the role of creep-induced stretching of a degraded aortic wall deserves a closer look because of its implications to the natural history of aneurysmal dilation. The goal of this work was to assess what role, if any, that passive creep plays in the stretching of porcine aortic wall by building upon previously reported work (Dadgar et al., 1997; Dobrin and Mrkvicka, 1994; Miskolczi et al., 1997). We set to verify the premise that a harvested aortic wall when deprived of elastin and/or collagen, will become increasingly susceptible to passive plastic deformation akin to an AAA. A typical surgically repaired AAA is about 5 cm or more in midsection diameter, which is a dilation of 150% from the 2 cm diameter abdominal aorta. We studied whether such levels of passive stretch may be induced in a harvested porcine aorta under varying protocols of enzymatic degradation and cyclic creep loading.
2.
Methods
2.1.
Study population
This study was conducted using thoracic aorta harvested from ten adult pigs (weight 260–320 lbs). The morphology of porcine thoracic aorta is comparable to human abdominal aorta. Multiple, 8 mm long ring specimens were cut from each thoracic aorta. The specimens were subjected to one of five protocols: Untreated controls (UC; N = 23), complete elastin degradation (E; N = 10), partial elastin degradation (Ep ; N = 10), partial collagen degradation (Cp ; N = 22), and partial degradation of both elastin and collagen (Ep + Cp ; N = 3). While each of the aortic trees had multiple paired control (UC) specimens, they did not necessarily have specimens from all five protocols. For a given aortic tree, treatment protocols for specimen rings were sequential. Fig. 1 illustrates the location of specimens from various protocols in one representative porcine aortic tree. In this case, the protocol order for consecutive specimens were UC, UC, E, Cp , Cp , UC, UC, E, Cp , Cp .
2.2.
Specimen preparation
The arterial tree from the aortic arch to the iliac bifurcation was harvested from each of the ten adult pigs. The locations of circumferential specimens were marked on the digital photograph of each aortic tree. Specimens were cut from approximately the same regions for all ten aortic trees (Fig. 1). Each specimen was then placed in a 0.9% saline solution and stored in the refrigerator at 4 ◦ C before enzymatic treatment and/or mechanical testing. All specimens were acquired, treated, and tested within 7 days, a period in which significant changes to aortic tissue mechanical properties are unlikely to occur (Samila and Carter, 1981). Given the duration of enzymatic treatments (1–5 days) and testing (1 day), this time window was inevitable.
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Fig. 1 – Aortic Tree and Specimen Locations Representative porcine aortic tree with sequentially cut specimens in the thoracic region and the treatment groups they belong to. A dime is placed in the figure to provide a scale. The toothpicks landmark branch points along the aortic tree.
2.3.
Treatment protocols
Prior to the enzymatic treatment, all specimen rings including the controls were treated with 0.1% sodium dodecylsulfate (SDS) solution at 37 ◦ C for 24 h to degrade the smooth muscle cells (SMCs) from the arterial matrix without causing mechanical degradation to the load bearing proteins of the arterial wall (Allaire et al., 1994; Malone et al., 1984). Removal of the SMCs was done to minimize bacterial growth and tissue degradation that may occur during enzyme treatments and storage. Further, since the goal was to study the scope for passive stretch of the aortic wall material, the removal of SMCs was desirable anyway. Besides, unlike muscular arteries, the contractile function of SMCs are of little importance to the mechanical characteristics of aorta (Bader, 1963; Dobrin, 1983), where passive properties predominate. The specimen rings were subjected to elastin and/or collagen degradation protocols using the proteolytic enzymes, elastase and collagenase, respectively. Treatment protocols were borrowed from earlier reports on degradation of elastin (Boudghene et al., 1993; Dobrin and Anidjar, 1991; Dobrin et al., 1984; Dobrin and Canfield, 1984; Dobrin and Mrkvicka, 1994) and collagen (Dadgar et al., 1997). Specimen rings were exposed to purified elastase and/or collagenase (Worthington Biochemical, Lakewood, NJ) in Tris buffered saline with pH 8.0 at 25 ◦ C. Table 1 shows the enzyme concentrations and treatment times used for all the four protocols. Different combinations of treatment times and concentrations were used for partial degradation of both elastin and collagen, thus resulting in varying levels of degradation. Selected specimen rings were histologically studied for elastin and collagen content. We noted that the treatment time was significantly more effective than enzyme concentration in controlling the level of digestion, and hence this was the main parameter that was varied.
2.4.
Mechanical loading protocol
The direct approach to studying the feasibility of developing an in vitro AAA would be to subject long tubular porcine aorta specimens (of at least 10 cm length) to proteolytic enzymes and then to high cyclic luminal pressures, while recording
Table 1 – Enzymatic treatment protocols Group
Conc. (Units/ml); time (h)
E (N = 10) Ep (N = 10) Cp (N = 22)
70; 6 (N = 10) 70; 2 (N = 10) 500; 24 (N = 8) 500; 12 (N = 2) 500; 2 (N = 3) 250; 6 (N = 8) 250; 2 (N = 1) 70; 2 + 500; 2 (N = 1) 70; 2 + 500; 2 (N = 1) 70; 2 + 500; 2 (N = 1)
Ep + Cp (N = 3)
Enzyme concentration and treatment times used for different levels of digestion of elastin and/or collagen. In all cases, Tris buffered saline with pH 8.0 at 25 ◦ C was used.
the resulting dilation. However, in order to increase the number of specimens tested under each treatment protocol, we investigated the maximum plastic strain inducible in rectangular, circumferentially oriented strips of the aortic wall before catastrophic failure occurs — an intrinsic property of the aortic wall. The maximum inducible plastic strain is likely to be the upper limit for the dilation inducible in a tubular aorta. If a plastic strain of 150% or more could be induced on the circumferential strips, then the development of an in vitro model would be a real possibility with just the practicalities to be worked out. Alternatively, if it is far lower than 150%, then developing an in vitro AAA is unlikely. This study therefore is a verification of the concept. The specimen rings were cut and flattened into rectangular flat strips for mechanical test protocols. Mechanical loading of the specimens was performed using an EMS Smart Test uniaxial extension tester with WINTEST controls and data acquisition system (EnduraTEC Systems Corporation, Minnetonka, MN). Preliminary testing on normal porcine aortic specimens was performed in order to determine the load increment and creep time that is likely to induce the maximal plastic strain, while also ensuring that the test is completed within a reasonable amount of time. The results of these preliminary studies showed that incrementally loading the specimens using small increments (1 or 2 N) and allowing the specimens to creep for 10 min, induced a greater plastic
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εpE =
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L −L Change in specimen length after enzyme treatment = E Specimen length before enzyme treatment L
εpL- max = εp- max =
δpL- max Load - induced maximum plastic displacement = Specimen length after enzyme treatment (before mechanical creep) LE
δpL- max + LE − L Total change in specimen length due to treatment and creep = Specimen length before enzyme treatment L Box I.
Fig. 2 – Mechanical Testing Grips Transparent grips used to clamp specimen strip for creep testing. A penny is placed in the figure to provide a scale.
deformation than applying large load increments (15–20 N) and allowing the specimens to creep for over 2 h. For creep
Fig. 3 – Load and Displacement vs. Time The controlled load-time curve and the measured displacement curve during an incremental cyclic creep test of a representative specimen.
loading, each specimen was placed between two specially designed tissue clamps (Fig. 2) and gripped by the Test machine. The specimens were sprinkled with 0.9% saline during the experiment in order to keep them hydrated and tested at room temperature. The initial length was established by extending the clamped specimen just until it was taut and an insignificantly small load of 0.05 N – resolution of load cell – was registered. The post-treatment specimen length (LE ), width, and thickness were then measured and recorded. For
for all subsequent cycles until failure. Based on the measurements, the plastic strains due to enzyme treatment alone (εpE ), loading alone (εpL- max ) and both together (εp- max ) were determined for each specimen (Box I). where L is the original pre-treatment inter-clamp length of the specimen determined by digitizing the images taken prior to treatment.
2.5.
Histology
the first creep cycle, the specimen was loaded to 2 N, allowed to creep for 10 min and then returned to the original length. During each creep phase, the displacement was recorded at three stages– 0, 5, and 10 min after the cycle’s load amplitude was reached. Continual visual observation of the displacement showed that by 10 min, the displacement rate had diminished significantly, although did not reach zero. Following creep, the specimen was returned to zero displacement, and then extended slowly until the load-cell registered its first non-zero reading of 0.05 N – its resolution – and the displacement measured. This is the zero-load displacement induced by the creep cycle (δpL ). For subsequent cycles, the target load amplitude was increased at 1 N increments from 2–10 N and then by 2 N increments for loads greater than 10 N until specimen failure occurred. Displacement measurements were repeated during and after each creep cycle. The load-induced maximum plastic displacement (δpL- max ) was defined as the δpL just before specimen failure occurred. Fig. 3 shows the load versus time and displacement versus time curves for a representative specimen. The measurements were repeated
Histological staining was performed to study the effect of elastase and collagenase on elastin and collagen in the aortic wall. For each specimen ring subjected to one of UC, E or C treatment protocols, an adjacent ring of 2–4 cm length was also cut from the aortic tree, subjected to an identical treatment protocol as their adjoining companion, fixed in a 10% formalin solution, and then embedded in paraffin. These companion rings were used for histological staining. Five micron (5 um) cross-sectional slices were then made across the length of the specimens and appropriate staining was performed in order to examine the extent of elastin and collagen degradation. Previous studies (Anidjar et al., 1992; Rizzo et al., 1989; Samila and Carter, 1981) have shown that the Verhoeff’s and van Gieson’s stains are appropriate for elastin and collagen, respectively, under light microscope conditions, and thus these two staining techniques were utilized. Once proper staining and had been completed, the microstructure was examined using an Olympus BX-51 bright light microscope with a viewing range from 4-60x magnification.
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Fig. 4 – Histology (a) Entire aortic cross-section with sodium dodecylsulfate (SDS) treatment only (black fibers elastin, purple fibers collagen); (b) Entire aortic cross-section after being treated for partial elastin degradation. (notice the removal of elastin fibers in the media); (c) Aortic cross-section treated for complete elastin degradation (notice how all black fibers are gone); (d) Aortic cross-section with media (left part of figure) and adventitia (right part of figure) untreated; (e) Aortic cross-section with media (left part of figure) and adventitia (right part of figure) treated for partial collagen degradation (Notice the coarseness of this figure as compared to d, it appears that much of the collagen has been removed, also notice the lack of structure in the adventitia).
3.
Results
Histological staining showed that the elastase treatment protocols for partial and complete elastin digestion were quite effective (see Fig. 4b,c). The collagenase treatment protocols showed partial collagen digestion, but remained unclear since collagen was not clearly isolated in the light microscope images of van Giessen’s stains. Additional support of collagen degradation exists in the failure load decrease seen in those specimens treated with collagenase vs. the untreated control (4.0 ± 1.06 N for 24 h Cp vs. 12.21 ± 4.90 N for 12 h Cp vs 15.86 ± 6.08 N for UC). Fig. 4a–e show the aortic microstructure under various treatment protocols.
Fig. 4a–c show tissue sections with Verhoeff’s stain in order to visualize the elastin (black) fibers in the arterial wall. Fig. 4d–e show sections stained with van Giessen’s stain in order to contrast the collagen (wavy purple) fibers. It should be noted that 4a–e were taken at 10x magnification in order to visualize the entire aortic cross-sections. For E and C specimens, εPE- max was determined by comparing the zero-load specimen length to its pretreatment length. Overall, a total of 69 specimens were subjected to cyclic creep loading until failure. Table 2 shows the average plastic strains recorded for UC, E and Ep , Cp , and Ep + Cp groups and the two-tailed p-values for paired Student’s t test in comparison to the UC group. The thickness
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Table 2 – Plastic Strains Induced for each Specimen UC
E
Ep
Cp
Ep + Cp
N
23
10
10
22
3
εpE (%)
– 17.7 ± 6.1 17.7 ± 6.1
22.9 ± 17.1# 22.5 ± 7.1 48.2 ± 17.6 *
15.0 ± 7.8 * 23.6 ± 10.5 41.6 ± 11.1 #
4.0 ± 9.0$ 17.5 ± 5.2 22.2 ± 12.8
15.9 ± 14.0 28.2 ± 3.3 48.9 ± 21.6
εpL- max (%) εp- max (%)
Plastic strains induced by enzyme treatment (εpE- max ), cyclic mechanical creep (εpL- max ) and the total (εp- max ) for various treatment groups. The groups found to be different from UC with statistical significance based on paired two-tailed Student’s t test are denoted by # (p < 0.0005), $ (p < 0.05) and ∗(p < 0.005). Where possible, UC specimens on either side of the treated specimen were averaged and used as the paired control for statistical comparisons to compensate for longitudinal variation in properties. An unpaired comparison resulted in the same findings for all groups except Ep+Cp which was significantly different (p < 0.00001). Although Ep and Cp specimens had various treatment times and concentrations, they were compared with their respective paired controls during statistical analyses. Additionally, one sample Student’s t test were performed to determine if the plastic strain due to the various enzyme treatments were statsitcally different than the UC group (where εPE- max = 0).
Table 3 – Thickness data for specimens
N Thickness (mm)
UC
E
Ep
Cp
Ep + C p
23 1.56 ± 0.37
10 1.24 ± 0.44$
10 1.07 ± 0.41*
22 1.06 ± 0.48#
3 1.19 ± 0.22
Thickness data for various treatment groups. The groups found to be different from UC with statistical ignificance are denoted by # (p < 0.0005), $ (p < 0.05) and ∗(p < 0.005).
Fig. 5 – Load and Displacement vs. Time (2 h creep) The controlled load-time curve and measured displacement curve for a 2 h incremental cyclic creep test of a representative specimen. Notice in the displacement vs. time curve that the creep curve begins to asymptote relatively quickly, reaching nearly 60% of the total displacement in the first 20 min.
of each specimen strip was also recorded (Table 3) in order to determine if enzyme treatment caused any significant changes.
4.
Discussion
Development of an in vitro AAA model can be valuable in conducting controlled experiments to study some aspects of this disease and in endovascular graft design. Some have speculated that mechanical stretch after selective enzymolysis can induce aneurysm-like dilations in a normal large animal aorta (Dadgar et al., 1997). We studied the
feasibility of developing such an in vitro AAA model. We adopted previously reported enzymolysis treatment protocols for our study, selectively degraded elastin and/or collagen in specimen strips, subjected them to cyclic creep tests of increasing load amplitude and recorded their plastic strain prior to failure. We found that the maximum plastic strain that may be induced in a porcine aortic tissue without tearing it apart was 45% when exposed to elastase. Alternatively, when exposed to collagenase, or when left untreated, specimens stretched plastically by less than 20%. Exposure to elastase followed by collagenase did not significantly alter the plastic stretch from exposure to elastase alone suggesting that the effect of these enzymes is not additive. One specimen in the E group registered the maximum plastic stretch (εpE- max , εpL- max , εp- max = 54.6%, 15.5%, 78.6%, respectively). Since none of the inducible plastic strains are close to the 150% sought, it appears that development of a usable in vitro AAA model is infeasible by this methodology. It has to be noted however that the low sample sizes of Cp , and Ep + Cp groups warrant some caution in interpreting their implications. By extension, these findings also suggest that aneurysm development is unlikely to be the result of passive stretching of a weakened wall, but rather predominated by growth and remodeling. This is also the first study to explicitly study the level of susceptibility to permanent stretching in the aortic wall. We have reported the level of stretch that may be created by enzymatic digestion alone, loading alone and both. This data may form the basis for attempts at developing mathematical models of vessel wall changes effected by hemodynamic perturbations. We also addressed the role of loading time on inducible plastic strain by performing an additional study in which specimens were tested in exactly the same manner as above except the loading time was increased to two hours between load increment intervals. In this study specimens
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Table 4 – Strain for 10 min vs. 2 h creep Load (N)
Creep Strain (mm/mm)
4 5 6 8
10 min Creep Test 0.09 0.11 0.11 0.13
2 h Creep Test 0.056 0.055 0.049 0.052
Total creep strain for two representative aortic specimens, one specimen tested using 10 min of creep and one specimen tested using 2 h of creep. The trend of increased strain in the 10 min creep experiments was seen throughout the experiments.
were subjected to UC, E, and Cp treatment protocols. The load induced plastic strain of each of these protocols (Mean ± SD) was 5.51% ± 0.91%, 10.17% ± 4.29%, 12.49% ± 0.21%, respectively. Comparing these strains to those specimens loaded for 10 min (Table 4) we see that increased load exposure does not increase the maximum inducible plastic strain. In addition, this set of experiments was also used to visualize the amount of creep that had occurred during the first 10 min of the creep cycle, which was found to be roughly 50% of the total creep that occurred over two hours (Fig. 5). As a result of this and the fact that the increased creep time did not cause an increase in the total plastic strain, it is felt that the 10 min cycles used in all other experiments gives a good representation of the maximum inducible plastic strain due to passive loading. Separating the plastic strains induced by enzyme treatment from that by creep loading provided some useful insights into the relative roles of elastin and collagen in the susceptibility of aortic tissue to plastic stretch. Enzymatic elastin degradation caused increased treatmentinduced plastic strain, while collagen degradation caused very little. This was also observed by Dobrin (1989), who reported qualitatively that the tubular artery “dilated” when exposed to elastase, but not to collagenase. This phenomenon may be attributed to the nature of elastin-collagen microstructure. The removal of taut elastin fibers is only likely to result in a transfer of load to the collagen fibers, which are tortuous when unloaded, thus having to plastically stretch until they too become taut enough for load bearing. The aortic tissue’s susceptibility to load-induced plastic strain was not significantly affected by elastin or collagen degradation. Dadgar et al. (1997) speculated that a weakened artery is likely to be more susceptible to load-induced plastic strain, but our results suggest that may not be the case. Our choice of porcine thoracic aorta is owing to its similarity in size to infrarenal human abdominal aorta. It is reasonable to wonder if the porcine infrarenal abdominal aorta is more susceptible to plastic stretch than its thoracic aorta. During this study, four specimens were harvested from the infrarenal portion of the aorta and treated with the UC and E protocols. We found that the normal specimens had an average total plastic strain of 32% ± 6% while the single specimen treated for complete elastin degradation had a total plastic strain of 42.7%. The normal specimens had a slightly increased plastic strain as compared to those specimens taken from the thoracic aorta, whereas the single specimen treated with elastase had an overall plastic strain very similar
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to the thoracic aorta specimens. Since this increase is very small compared to the overall 150% strain needed to create aneurysm conditions, we feel that our findings would remain the same if abdominal aortic specimens were used. The creep test parameters such as creep time and load increment were chosen based on preliminary studies. We found that εpL- max was not affected when the creep time was increased to two hours per cycle. Preliminary studies also indicated that a low load-increment is more likely to induce greater plastic stretch, although not statistically so. A low load-increment is desirable anyway so as to increase the resolution of our δpL measurements. We believe that our loading protocol is likely to have resulted in the maximum possible plastic stretch that the porcine aorta can sustain prior to failure. To the best of our knowledge, this study is the first to perform controlled experiments to investigate the effect of selective enzymolysis on the susceptibility of aortic tissue to plastic strain. Experimental limitations associated with this work should be addressed in the future. The effectiveness of enzyme treatments in selectively degrading elastin or collagen was not quantatively confirmed, especially for collagenase. Histological verification of elastin degradation is quite clear on stained slides, with E specimens exhibiting no remaining elastin in the aortic wall (Fig. 4c) and those treated for partial degradation displaying some elastin in the center of the aortic wall (Fig. 4b). When exposed to collagenase however, the collagen content appears to have decreased for some specimens and unclear in others. This is further supported by the fact that the failure load for collagenase treated specimens is significantly decreased compared to the control group, suggesting some collagen depletion is occuring. It must be noted that the current collagenase treatment was performed at room temperature rather than at body temperature (Dadgar et al., 1997) and this may have limited its potency. Alternatively, it is possible that collagen may still be present in the aortic wall even after being treated with collagenase (as seen histologically), but may not be coherent enough to be of any mechanical consequence. We know that human AAA have a significant amount of collagen remaining in the aortic wall despite decreased failure properties (Menashi et al., 1987; Rizzo et al., 1989), a possible explanation for the results we see here. If the collagen was not completely depleted in these tissues we would expect the tissue to be able to undergo a greater plastic strain than a tissue with no collagen because the latter would fail prematurely. Thus, the difficulty with collagen digestion may have overestimated the plastic strain, not underestimated it. A second limitation is the low resolution in displacement measurements owing to lack of an automated displacement data acquisition system during the time of these experiments. The accuracy of εpE- max may be affected slightly by the predicted pre-treatment length of the specimen since some specimen-to-specimen differences in clamping is likely to exist. A final limitation is that the specimens were not allowed to creep in the reverse direction once returned to its zero load length. Recent experiments show that a specimen returned from a creep cycle to zero load will continue to reduce in length while at that zero-load (reverse creeping). The plastic strain was typically measured within about 1 min of reaching zero load length. This will lead to a slight over estimation (about 3% according to recent preliminary studies) of the load-induced plastic strain. We do
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not believe these limitations will affect the overall finding since an overestimation would be a worst case scenario for the plastic strain. Thus the conclusion that the creation of an in vitro AAA via enzymatic treatments and mechanical loading is not feasible should remain valid. An in vitro AAA model can aid in many aspects of endovascular treatment of AAA. We found no evidence to support the feasibility of developing an in vitro AAA model by passively degrading the load bearing fibers of the aortic wall and subjecting these tissues to cyclic creep. Future attempts at developing an in vitro model may need to consider alternative approaches such as experimentation over longer durations (weeks and months, not days), in which the cells in the vascular wall are kept alive in a bioreactor environment and slowly exposed to proteolytic enzymes while subjected to gradually increasing cyclic pressure. Incorporation of tissue remodeling into an in vitro experimental setup is inevitable for development of in vitro AAA. We also found little evidence to support the postulate that passive creep may play an important role in aneurysmal dilation of a degraded aortic wall. REFERENCES
Allaire, E., Guettier, C., Bruneval, P., Plissonnier, D., Michel, J.B., 1994. Cell-free arterial grafts: Morphologic characteristics of aortic isografts, allografts, and xenografts in rats. J. Vasc. Surg. 19, 446–456. Andrews, E.J., White, W.J., Bullock, L.P., 1975. Spontaneous aortic aneurysms in blotchy mice. Am. J. Pathol 78, 199–210. Anidjar, S., Dobrin, P.B., Eichorst, M., Graham, G.P., Chejfec, G., 1992. Correlation of inflammatory infiltrate with the enlargement of experimental aortic aneurysms. J. Vasc. Surg. 16, 139–147. Anidjar, S., Salzmann, J.L., Gentric, D., Lagneau, P., Camilleri, J.P., Michel, J.B., 1990. Elastase-induced experimental aneurysms in rats. Circulation 82, 973–981. Bader, H., 1963. Handbook of Physiology, Part II (Circulation). In: Book Handbook of Physiology, Part II (Circulation), American Physiological Society, pp. 875–876. Baxter, B.T., Davis, V.A., Minion, D.J., Wang, Y.P., Lynch, T.G., McManus, B.M., 1994. Abdominal aortic aneurysms are associated with altered matrix proteins of the nonaneurysmal aortic segments. J. Vasc. Surg. 19, 797–802. discussion 803. Baxter, T.B., Halloran, B.G., 1994. Aneurysms: New Findings and Treatments. In: Book Aneurysms: New Findings and Treatments, Appleton and Lange, pp. 25–34. Boudghene, F., Anidjar, S., Allaire, E., Osborne-Pellegrin, M., Bigot, J.M., Michel, J.B., 1993. Endovascular grafting in elastaseinduced experimental aortic aneurysms in dogs: Feasibility and preliminary results. J. Vasc. Interv. Radiol. 4, 497–504.
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Dadgar, L., Marois, Y., Deng, X., Guidoin, R., 1997. Arterial wall mechanical characteristics after treatment in collagenase: An in vitro aneurysm model. Clin. Invest. Med. 20, 25–34. Davis, N.P., Han, H.C., Wayman, B., Vito, R., 2005. Sustained axial loading lengthens arteries in organ culture. Ann. Biomed. Eng. 33, 867–877. Dobrin, P.B., 1983. Handbook of Physiology, Pat I (Peripheral Circulation and Organ Blood Flow. In: Book Handbook of Physiology, Williams and Wilkins, pp. 65–102. Dobrin, P.B., 1989. Pathophysiology and pathogenesis of aortic aneurysms. Current concepts. Surg. Clin. North. Am. 69, 687–703. Dobrin, P.B., Anidjar, S., 1991. Pathophysiology of arterial aneurysms. Arch. Mal, Coeur. Vaiss 84 (3), 57–62. Dobrin, P.B., Baker, W.H., Gley, W.C., 1984. Elastolytic and collagenolytic studies of arteries. Implications for the mechanical properties of aneurysms. Arch. Surg. 119, 405–409. Dobrin, P.B., Canfield, T.R., 1984. Elastase, collagenase, and the biaxial elastic properties of dog carotid artery. Am. J. Physiol. 247, H124–H131. Dobrin, P.B., Mrkvicka, R., 1994. Failure of elastin or collagen as possible critical connective tissue alterations underlying aneurysmal dilatation. Cardiovasc. Surg. 2, 484–488. Gertz, S.D., Kurgan, A., Eisenberg, D., 1988. Aneurysm of the rabbit common carotid artery induced by periarterial application of calcium chloride in vivo. J. Clin. Invest. 81, 649–656. Keen, R.R., Dobrin, P.B., 2000. Development of Aneurysms. In: Book Development of Aneurysms, Landes Bioscience and Eureka.com. MacSweeney, S.T., Powell, J.T., Greenhalgh, R.M., 1994. Pathogenesis of abdominal aortic aneurysm. Br. J. Surg. 81, 935–941. Malone, J.M., Brendel, K., Duhamel, R.C., Reinert, R.L., 1984. Detergent-extracted small-diameter vascular prostheses. J. Vasc. Surg. 1, 181–191. Marinov, G.R., Marois, Y., Paris, E., Roby, P., Formichi, M., Douville, Y., Guidoin, R., 1997. Can the infusion of elastase in the abdominal aorta of the Yucatan miniature swine consistently produce experimental aneurysms?. J. Invest. Surg. 10, 129–150. Menashi, S., Campa, J.S., Greenhalgh, R.M., Powell, J.T., 1987. Collagen in abdominal aortic aneurysm: Typing, content, and degradation. J. Vasc. Surg. 6, 578–582. Miskolczi, L., Guterman, L.R., Flaherty, J.D., Szikora, I., Hopkins, L.N., 1997. Rapid saccular aneurysm induction by elastase application in vitro. Neurosurgery 41, 220–228. discussion 228–229. Rizzo, R.J., McCarthy, W.J., Dixit, S.N., Lilly, M.P., Shively, V.P., Flinn, W.R., Yao, J.S., 1989. Collagen types and matrix protein content in human abdominal aortic aneurysms. J. Vasc. Surg. 10, 365–373. Samila, Z.J., Carter, S.A., 1981. The effect of age on the unfolding of elastin lamellae and collagen fibers with stretch in human carotid arteries. Can. J. Physiol. Pharmacol. 59, 1050–1057. Yang, X.J., Li, L., Wu, Z.X., 2007. A novel arterial pouch model of saccular aneurysm by concomitant elastase and collagenase digestion. J. Zhejiang Univ. Sci. B 8, 697–703.
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available at www.sciencedirect.com
journal homepage: www.elsevier.com/locate/jmbbm
Research paper
The effect of proteolytic treatment on plastic deformation of porcine aortic tissue Jarin A. Kratzberg, Patricia J. Walker, Elizabeth Rikkers, Madhavan L. Raghavan ∗ Department of Biomedical Engineering, University of Iowa, Iowa City, IA, United States
A R T I C L E
I N F O
A B S T R A C T
Article history:
The objective of this work was to assess whether selective proteolysis of elastin and/or col-
Received 25 May 2007
lagen in a porcine aorta followed by mechanical creep loading would result in an aneurysm-
Received in revised form
like permanent tissue stretch. The underlying motivations were to (1) test the feasibility of
2 April 2008
developing an in vitro abdominal aortic aneurysm (AAA) model, and (2) understand what
Accepted 4 April 2008
role, if any, that passive creep-induced stretching plays in aneurysmal dilation. Multiple
Published online 23 April 2008
circumferentially oriented flat specimen strips were cut from the porcine thoracic aorta of ten adult pigs. Specimens were subjected to one of six treatment protocols: Untreated
Keywords:
controls (UC; N = 23), complete elastin degradation (E; N = 10), partial elastin degrada-
Aneurysm
tion (Ep ; N = 10), partial collagen degradation (Cp ; N = 22), and partial degradation of both
AAA
elastin and collagen (Ep + Cp ; N = 3). All specimens were then subjected to cyclic creep (10
Elastin
min/cycle) with increasing load amplitude until failure. The zero-load strain prior to the
Collagen
creep cycle where failure occurred was defined as load-induced plastic strain. The plastic
Enzyme degradation
strain induced by treatment alone, creep loading alone and the total was determined for all specimens. The total plastic strain was significantly greater for E (mean ± SD = 48.2 ±17.6, p < 0.005), Ep (41.6 ± 11.1, p < 0.0005), but not for Ep + Cp (48.9 ± 21.6, p = 0.17) and Cp (22.2 ± 12.8, p = 0.14) compared to UC (17.7 ± 6.1). Of the total plastic strain, treatmentinduced plastic strain was high for those specimens subjected to partial or total elastin degradation (E, Ep , Ep + Cp ), but not for those where elastin was intact (Cp ). However, loadinduced plastic strain in the treated specimens was not different in any of the treated groups compared to controls. Maximum total plastic strain of 78.6% was induced in one porcine aortic tissue from the E group. Even this is far lower than what would be needed for creating a realistic in vitro AAA model. Our findings do not support the feasibility of developing an in vitro AAA by enzymolysis followed by passive stretching of the aorta. The findings also suggest that AAA formation is unlikely to be a passive creep-induced stretching of a proteolytically degraded aortic wall as conventional thinking may suggest, but rather may be predominantly due to growth and remodeling of the aortic wall. c 2008 Elsevier Ltd. All rights reserved.
1.
Background
Abdominal aortic aneurysm (AAA) is a localized dilation of the abdominal aorta. AAA is characterized by a thinning
media with decreased elastin, thickening of the adventitia (outermost layer of the aorta) and neointima (Baxter and Halloran, 1994), and an increase in proteolytic enzyme (collaganase and elastase) synthesis. AAA pathogenesis is
∗ Corresponding address: Department of Biomedical Engineering, 1422 Seamans Center (SC), University of Iowa, Iowa City, IA 52242-1527, United States. Tel.: +1 319 335 5704; fax: +1 319 335 5631. E-mail address: [email protected] (M.L. Raghavan). c 2008 Elsevier Ltd. All rights reserved. 1751-6161/$ - see front matter doi:10.1016/j.jmbbm.2008.04.001
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thought to be triggered by enzymatic degradation of elastin combined with repetitive pressurization of the aorta and remodeling of collagen (Baxter et al., 1994; Dobrin, 1983; Keen and Dobrin, 2000; MacSweeney et al., 1994). A reliable animal model of AAA can vastly aid research toward understanding this disease. Researchers have been successful in developing in vivo aneurysm models in small animals such as rats, mice, and rabbits (Andrews et al., 1975; Anidjar et al., 1990; Gertz et al., 1988; Yang et al., 2007) by perfusing the abdominal aorta with elastase – a proteolytic enzyme that can selectively digest elastin. Such small animal models have been valuable in studying genetic and histological aspects of AAA, but not biomechanical characteristics or graft implant function owing to their miniature size. There has been only one report of similar success in large animals. Boudghene et al. (1993) reported successfully inducing a small dilation in dog abdominal aorta in vivo using the elastase perfusion method for the purpose of testing endovascular grafts. Similar attempts in other animals have not been successful (Marinov et al., 1997). The ability to create reliable large animal models of AAA remains elusive. Dadgar et al. (1997) proposed the development of an in vitro AAA model using harvested porcine thoracic or abdominal aorta. In this preliminary study, they showed that the excised porcine aorta is weakened by exposure to collagenase – proteolytic enzyme that selectively digests collagen – and that the level of this weakening may be controlled by exposure time. They hypothesized that a weakened aorta may then be biomechanically stretched in a laboratory to induce an aneurysm-like dilation, thus creating a dilated aorta with the morphological characteristics of a human AAA. Earlier, in an in vitro study, Dobrin and Mrkvicka (1994) reported that harvested tubular human iliac artery dilated in vitro when exposed to elastase, but ruptured when exposed to collagenase under physiologic pressures. Miskolczi et al. (1997) reported that porcine carotid artery dilated slightly at its bifurcation when exposed to elastase and subjected to pulsatile pressure in a saline bath. Neither of these studies (Dobrin and Mrkvicka, 1994; Miskolczi et al., 1997) reported the level of zero-load dilation achieved, however small they were. More recently Davis et al. (2005) reported plastic deformation in porcine common carotid arteries due to axial loading. This is an important discovery, as it shows that passive loading of an artery can cause permanent deformation. However, the feasibility of developing an in vitro AAA model through the combination of passive loading and enzymatic treatment of healthy aortic tissue has not been addressed. An in vitro AAA model would be a valuable tool for studying the effects of dilation on tissue properties, but it is unlikely to be of value in studying histological or genetic aspects of this disease since it does not accommodate the active remodeling that is undoubtedly taking place in vivo. Perhaps the most valuable utility of an in vitro AAA model is in testing new designs of endovascular grafts. Graft delivery designs, graft-to-aorta attachment designs and stent architecture designs may all be studied in highly controlled environments with in vitro AAA models. Further, fatigue testing of an endovascular graft implanted in an in vitro AAA would be substantially more realistic than current
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approaches where compliant cylindrical tubes are used. An in vitro AAA model can also be valuable in predicting shortterm implant performance such as post-procedural endoleak, although it is unlikely to be a good model for long-term complications. Secondly, the role of creep-induced stretching of a degraded aortic wall deserves a closer look because of its implications to the natural history of aneurysmal dilation. The goal of this work was to assess what role, if any, that passive creep plays in the stretching of porcine aortic wall by building upon previously reported work (Dadgar et al., 1997; Dobrin and Mrkvicka, 1994; Miskolczi et al., 1997). We set to verify the premise that a harvested aortic wall when deprived of elastin and/or collagen, will become increasingly susceptible to passive plastic deformation akin to an AAA. A typical surgically repaired AAA is about 5 cm or more in midsection diameter, which is a dilation of 150% from the 2 cm diameter abdominal aorta. We studied whether such levels of passive stretch may be induced in a harvested porcine aorta under varying protocols of enzymatic degradation and cyclic creep loading.
2.
Methods
2.1.
Study population
This study was conducted using thoracic aorta harvested from ten adult pigs (weight 260–320 lbs). The morphology of porcine thoracic aorta is comparable to human abdominal aorta. Multiple, 8 mm long ring specimens were cut from each thoracic aorta. The specimens were subjected to one of five protocols: Untreated controls (UC; N = 23), complete elastin degradation (E; N = 10), partial elastin degradation (Ep ; N = 10), partial collagen degradation (Cp ; N = 22), and partial degradation of both elastin and collagen (Ep + Cp ; N = 3). While each of the aortic trees had multiple paired control (UC) specimens, they did not necessarily have specimens from all five protocols. For a given aortic tree, treatment protocols for specimen rings were sequential. Fig. 1 illustrates the location of specimens from various protocols in one representative porcine aortic tree. In this case, the protocol order for consecutive specimens were UC, UC, E, Cp , Cp , UC, UC, E, Cp , Cp .
2.2.
Specimen preparation
The arterial tree from the aortic arch to the iliac bifurcation was harvested from each of the ten adult pigs. The locations of circumferential specimens were marked on the digital photograph of each aortic tree. Specimens were cut from approximately the same regions for all ten aortic trees (Fig. 1). Each specimen was then placed in a 0.9% saline solution and stored in the refrigerator at 4 ◦ C before enzymatic treatment and/or mechanical testing. All specimens were acquired, treated, and tested within 7 days, a period in which significant changes to aortic tissue mechanical properties are unlikely to occur (Samila and Carter, 1981). Given the duration of enzymatic treatments (1–5 days) and testing (1 day), this time window was inevitable.
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Fig. 1 – Aortic Tree and Specimen Locations Representative porcine aortic tree with sequentially cut specimens in the thoracic region and the treatment groups they belong to. A dime is placed in the figure to provide a scale. The toothpicks landmark branch points along the aortic tree.
2.3.
Treatment protocols
Prior to the enzymatic treatment, all specimen rings including the controls were treated with 0.1% sodium dodecylsulfate (SDS) solution at 37 ◦ C for 24 h to degrade the smooth muscle cells (SMCs) from the arterial matrix without causing mechanical degradation to the load bearing proteins of the arterial wall (Allaire et al., 1994; Malone et al., 1984). Removal of the SMCs was done to minimize bacterial growth and tissue degradation that may occur during enzyme treatments and storage. Further, since the goal was to study the scope for passive stretch of the aortic wall material, the removal of SMCs was desirable anyway. Besides, unlike muscular arteries, the contractile function of SMCs are of little importance to the mechanical characteristics of aorta (Bader, 1963; Dobrin, 1983), where passive properties predominate. The specimen rings were subjected to elastin and/or collagen degradation protocols using the proteolytic enzymes, elastase and collagenase, respectively. Treatment protocols were borrowed from earlier reports on degradation of elastin (Boudghene et al., 1993; Dobrin and Anidjar, 1991; Dobrin et al., 1984; Dobrin and Canfield, 1984; Dobrin and Mrkvicka, 1994) and collagen (Dadgar et al., 1997). Specimen rings were exposed to purified elastase and/or collagenase (Worthington Biochemical, Lakewood, NJ) in Tris buffered saline with pH 8.0 at 25 ◦ C. Table 1 shows the enzyme concentrations and treatment times used for all the four protocols. Different combinations of treatment times and concentrations were used for partial degradation of both elastin and collagen, thus resulting in varying levels of degradation. Selected specimen rings were histologically studied for elastin and collagen content. We noted that the treatment time was significantly more effective than enzyme concentration in controlling the level of digestion, and hence this was the main parameter that was varied.
2.4.
Mechanical loading protocol
The direct approach to studying the feasibility of developing an in vitro AAA would be to subject long tubular porcine aorta specimens (of at least 10 cm length) to proteolytic enzymes and then to high cyclic luminal pressures, while recording
Table 1 – Enzymatic treatment protocols Group
Conc. (Units/ml); time (h)
E (N = 10) Ep (N = 10) Cp (N = 22)
70; 6 (N = 10) 70; 2 (N = 10) 500; 24 (N = 8) 500; 12 (N = 2) 500; 2 (N = 3) 250; 6 (N = 8) 250; 2 (N = 1) 70; 2 + 500; 2 (N = 1) 70; 2 + 500; 2 (N = 1) 70; 2 + 500; 2 (N = 1)
Ep + Cp (N = 3)
Enzyme concentration and treatment times used for different levels of digestion of elastin and/or collagen. In all cases, Tris buffered saline with pH 8.0 at 25 ◦ C was used.
the resulting dilation. However, in order to increase the number of specimens tested under each treatment protocol, we investigated the maximum plastic strain inducible in rectangular, circumferentially oriented strips of the aortic wall before catastrophic failure occurs — an intrinsic property of the aortic wall. The maximum inducible plastic strain is likely to be the upper limit for the dilation inducible in a tubular aorta. If a plastic strain of 150% or more could be induced on the circumferential strips, then the development of an in vitro model would be a real possibility with just the practicalities to be worked out. Alternatively, if it is far lower than 150%, then developing an in vitro AAA is unlikely. This study therefore is a verification of the concept. The specimen rings were cut and flattened into rectangular flat strips for mechanical test protocols. Mechanical loading of the specimens was performed using an EMS Smart Test uniaxial extension tester with WINTEST controls and data acquisition system (EnduraTEC Systems Corporation, Minnetonka, MN). Preliminary testing on normal porcine aortic specimens was performed in order to determine the load increment and creep time that is likely to induce the maximal plastic strain, while also ensuring that the test is completed within a reasonable amount of time. The results of these preliminary studies showed that incrementally loading the specimens using small increments (1 or 2 N) and allowing the specimens to creep for 10 min, induced a greater plastic
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εpE =
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L −L Change in specimen length after enzyme treatment = E Specimen length before enzyme treatment L
εpL- max = εp- max =
δpL- max Load - induced maximum plastic displacement = Specimen length after enzyme treatment (before mechanical creep) LE
δpL- max + LE − L Total change in specimen length due to treatment and creep = Specimen length before enzyme treatment L Box I.
Fig. 2 – Mechanical Testing Grips Transparent grips used to clamp specimen strip for creep testing. A penny is placed in the figure to provide a scale.
deformation than applying large load increments (15–20 N) and allowing the specimens to creep for over 2 h. For creep
Fig. 3 – Load and Displacement vs. Time The controlled load-time curve and the measured displacement curve during an incremental cyclic creep test of a representative specimen.
loading, each specimen was placed between two specially designed tissue clamps (Fig. 2) and gripped by the Test machine. The specimens were sprinkled with 0.9% saline during the experiment in order to keep them hydrated and tested at room temperature. The initial length was established by extending the clamped specimen just until it was taut and an insignificantly small load of 0.05 N – resolution of load cell – was registered. The post-treatment specimen length (LE ), width, and thickness were then measured and recorded. For
for all subsequent cycles until failure. Based on the measurements, the plastic strains due to enzyme treatment alone (εpE ), loading alone (εpL- max ) and both together (εp- max ) were determined for each specimen (Box I). where L is the original pre-treatment inter-clamp length of the specimen determined by digitizing the images taken prior to treatment.
2.5.
Histology
the first creep cycle, the specimen was loaded to 2 N, allowed to creep for 10 min and then returned to the original length. During each creep phase, the displacement was recorded at three stages– 0, 5, and 10 min after the cycle’s load amplitude was reached. Continual visual observation of the displacement showed that by 10 min, the displacement rate had diminished significantly, although did not reach zero. Following creep, the specimen was returned to zero displacement, and then extended slowly until the load-cell registered its first non-zero reading of 0.05 N – its resolution – and the displacement measured. This is the zero-load displacement induced by the creep cycle (δpL ). For subsequent cycles, the target load amplitude was increased at 1 N increments from 2–10 N and then by 2 N increments for loads greater than 10 N until specimen failure occurred. Displacement measurements were repeated during and after each creep cycle. The load-induced maximum plastic displacement (δpL- max ) was defined as the δpL just before specimen failure occurred. Fig. 3 shows the load versus time and displacement versus time curves for a representative specimen. The measurements were repeated
Histological staining was performed to study the effect of elastase and collagenase on elastin and collagen in the aortic wall. For each specimen ring subjected to one of UC, E or C treatment protocols, an adjacent ring of 2–4 cm length was also cut from the aortic tree, subjected to an identical treatment protocol as their adjoining companion, fixed in a 10% formalin solution, and then embedded in paraffin. These companion rings were used for histological staining. Five micron (5 um) cross-sectional slices were then made across the length of the specimens and appropriate staining was performed in order to examine the extent of elastin and collagen degradation. Previous studies (Anidjar et al., 1992; Rizzo et al., 1989; Samila and Carter, 1981) have shown that the Verhoeff’s and van Gieson’s stains are appropriate for elastin and collagen, respectively, under light microscope conditions, and thus these two staining techniques were utilized. Once proper staining and had been completed, the microstructure was examined using an Olympus BX-51 bright light microscope with a viewing range from 4-60x magnification.
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Fig. 4 – Histology (a) Entire aortic cross-section with sodium dodecylsulfate (SDS) treatment only (black fibers elastin, purple fibers collagen); (b) Entire aortic cross-section after being treated for partial elastin degradation. (notice the removal of elastin fibers in the media); (c) Aortic cross-section treated for complete elastin degradation (notice how all black fibers are gone); (d) Aortic cross-section with media (left part of figure) and adventitia (right part of figure) untreated; (e) Aortic cross-section with media (left part of figure) and adventitia (right part of figure) treated for partial collagen degradation (Notice the coarseness of this figure as compared to d, it appears that much of the collagen has been removed, also notice the lack of structure in the adventitia).
3.
Results
Histological staining showed that the elastase treatment protocols for partial and complete elastin digestion were quite effective (see Fig. 4b,c). The collagenase treatment protocols showed partial collagen digestion, but remained unclear since collagen was not clearly isolated in the light microscope images of van Giessen’s stains. Additional support of collagen degradation exists in the failure load decrease seen in those specimens treated with collagenase vs. the untreated control (4.0 ± 1.06 N for 24 h Cp vs. 12.21 ± 4.90 N for 12 h Cp vs 15.86 ± 6.08 N for UC). Fig. 4a–e show the aortic microstructure under various treatment protocols.
Fig. 4a–c show tissue sections with Verhoeff’s stain in order to visualize the elastin (black) fibers in the arterial wall. Fig. 4d–e show sections stained with van Giessen’s stain in order to contrast the collagen (wavy purple) fibers. It should be noted that 4a–e were taken at 10x magnification in order to visualize the entire aortic cross-sections. For E and C specimens, εPE- max was determined by comparing the zero-load specimen length to its pretreatment length. Overall, a total of 69 specimens were subjected to cyclic creep loading until failure. Table 2 shows the average plastic strains recorded for UC, E and Ep , Cp , and Ep + Cp groups and the two-tailed p-values for paired Student’s t test in comparison to the UC group. The thickness
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Table 2 – Plastic Strains Induced for each Specimen UC
E
Ep
Cp
Ep + Cp
N
23
10
10
22
3
εpE (%)
– 17.7 ± 6.1 17.7 ± 6.1
22.9 ± 17.1# 22.5 ± 7.1 48.2 ± 17.6 *
15.0 ± 7.8 * 23.6 ± 10.5 41.6 ± 11.1 #
4.0 ± 9.0$ 17.5 ± 5.2 22.2 ± 12.8
15.9 ± 14.0 28.2 ± 3.3 48.9 ± 21.6
εpL- max (%) εp- max (%)
Plastic strains induced by enzyme treatment (εpE- max ), cyclic mechanical creep (εpL- max ) and the total (εp- max ) for various treatment groups. The groups found to be different from UC with statistical significance based on paired two-tailed Student’s t test are denoted by # (p < 0.0005), $ (p < 0.05) and ∗(p < 0.005). Where possible, UC specimens on either side of the treated specimen were averaged and used as the paired control for statistical comparisons to compensate for longitudinal variation in properties. An unpaired comparison resulted in the same findings for all groups except Ep+Cp which was significantly different (p < 0.00001). Although Ep and Cp specimens had various treatment times and concentrations, they were compared with their respective paired controls during statistical analyses. Additionally, one sample Student’s t test were performed to determine if the plastic strain due to the various enzyme treatments were statsitcally different than the UC group (where εPE- max = 0).
Table 3 – Thickness data for specimens
N Thickness (mm)
UC
E
Ep
Cp
Ep + C p
23 1.56 ± 0.37
10 1.24 ± 0.44$
10 1.07 ± 0.41*
22 1.06 ± 0.48#
3 1.19 ± 0.22
Thickness data for various treatment groups. The groups found to be different from UC with statistical ignificance are denoted by # (p < 0.0005), $ (p < 0.05) and ∗(p < 0.005).
Fig. 5 – Load and Displacement vs. Time (2 h creep) The controlled load-time curve and measured displacement curve for a 2 h incremental cyclic creep test of a representative specimen. Notice in the displacement vs. time curve that the creep curve begins to asymptote relatively quickly, reaching nearly 60% of the total displacement in the first 20 min.
of each specimen strip was also recorded (Table 3) in order to determine if enzyme treatment caused any significant changes.
4.
Discussion
Development of an in vitro AAA model can be valuable in conducting controlled experiments to study some aspects of this disease and in endovascular graft design. Some have speculated that mechanical stretch after selective enzymolysis can induce aneurysm-like dilations in a normal large animal aorta (Dadgar et al., 1997). We studied the
feasibility of developing such an in vitro AAA model. We adopted previously reported enzymolysis treatment protocols for our study, selectively degraded elastin and/or collagen in specimen strips, subjected them to cyclic creep tests of increasing load amplitude and recorded their plastic strain prior to failure. We found that the maximum plastic strain that may be induced in a porcine aortic tissue without tearing it apart was 45% when exposed to elastase. Alternatively, when exposed to collagenase, or when left untreated, specimens stretched plastically by less than 20%. Exposure to elastase followed by collagenase did not significantly alter the plastic stretch from exposure to elastase alone suggesting that the effect of these enzymes is not additive. One specimen in the E group registered the maximum plastic stretch (εpE- max , εpL- max , εp- max = 54.6%, 15.5%, 78.6%, respectively). Since none of the inducible plastic strains are close to the 150% sought, it appears that development of a usable in vitro AAA model is infeasible by this methodology. It has to be noted however that the low sample sizes of Cp , and Ep + Cp groups warrant some caution in interpreting their implications. By extension, these findings also suggest that aneurysm development is unlikely to be the result of passive stretching of a weakened wall, but rather predominated by growth and remodeling. This is also the first study to explicitly study the level of susceptibility to permanent stretching in the aortic wall. We have reported the level of stretch that may be created by enzymatic digestion alone, loading alone and both. This data may form the basis for attempts at developing mathematical models of vessel wall changes effected by hemodynamic perturbations. We also addressed the role of loading time on inducible plastic strain by performing an additional study in which specimens were tested in exactly the same manner as above except the loading time was increased to two hours between load increment intervals. In this study specimens
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Table 4 – Strain for 10 min vs. 2 h creep Load (N)
Creep Strain (mm/mm)
4 5 6 8
10 min Creep Test 0.09 0.11 0.11 0.13
2 h Creep Test 0.056 0.055 0.049 0.052
Total creep strain for two representative aortic specimens, one specimen tested using 10 min of creep and one specimen tested using 2 h of creep. The trend of increased strain in the 10 min creep experiments was seen throughout the experiments.
were subjected to UC, E, and Cp treatment protocols. The load induced plastic strain of each of these protocols (Mean ± SD) was 5.51% ± 0.91%, 10.17% ± 4.29%, 12.49% ± 0.21%, respectively. Comparing these strains to those specimens loaded for 10 min (Table 4) we see that increased load exposure does not increase the maximum inducible plastic strain. In addition, this set of experiments was also used to visualize the amount of creep that had occurred during the first 10 min of the creep cycle, which was found to be roughly 50% of the total creep that occurred over two hours (Fig. 5). As a result of this and the fact that the increased creep time did not cause an increase in the total plastic strain, it is felt that the 10 min cycles used in all other experiments gives a good representation of the maximum inducible plastic strain due to passive loading. Separating the plastic strains induced by enzyme treatment from that by creep loading provided some useful insights into the relative roles of elastin and collagen in the susceptibility of aortic tissue to plastic stretch. Enzymatic elastin degradation caused increased treatmentinduced plastic strain, while collagen degradation caused very little. This was also observed by Dobrin (1989), who reported qualitatively that the tubular artery “dilated” when exposed to elastase, but not to collagenase. This phenomenon may be attributed to the nature of elastin-collagen microstructure. The removal of taut elastin fibers is only likely to result in a transfer of load to the collagen fibers, which are tortuous when unloaded, thus having to plastically stretch until they too become taut enough for load bearing. The aortic tissue’s susceptibility to load-induced plastic strain was not significantly affected by elastin or collagen degradation. Dadgar et al. (1997) speculated that a weakened artery is likely to be more susceptible to load-induced plastic strain, but our results suggest that may not be the case. Our choice of porcine thoracic aorta is owing to its similarity in size to infrarenal human abdominal aorta. It is reasonable to wonder if the porcine infrarenal abdominal aorta is more susceptible to plastic stretch than its thoracic aorta. During this study, four specimens were harvested from the infrarenal portion of the aorta and treated with the UC and E protocols. We found that the normal specimens had an average total plastic strain of 32% ± 6% while the single specimen treated for complete elastin degradation had a total plastic strain of 42.7%. The normal specimens had a slightly increased plastic strain as compared to those specimens taken from the thoracic aorta, whereas the single specimen treated with elastase had an overall plastic strain very similar
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to the thoracic aorta specimens. Since this increase is very small compared to the overall 150% strain needed to create aneurysm conditions, we feel that our findings would remain the same if abdominal aortic specimens were used. The creep test parameters such as creep time and load increment were chosen based on preliminary studies. We found that εpL- max was not affected when the creep time was increased to two hours per cycle. Preliminary studies also indicated that a low load-increment is more likely to induce greater plastic stretch, although not statistically so. A low load-increment is desirable anyway so as to increase the resolution of our δpL measurements. We believe that our loading protocol is likely to have resulted in the maximum possible plastic stretch that the porcine aorta can sustain prior to failure. To the best of our knowledge, this study is the first to perform controlled experiments to investigate the effect of selective enzymolysis on the susceptibility of aortic tissue to plastic strain. Experimental limitations associated with this work should be addressed in the future. The effectiveness of enzyme treatments in selectively degrading elastin or collagen was not quantatively confirmed, especially for collagenase. Histological verification of elastin degradation is quite clear on stained slides, with E specimens exhibiting no remaining elastin in the aortic wall (Fig. 4c) and those treated for partial degradation displaying some elastin in the center of the aortic wall (Fig. 4b). When exposed to collagenase however, the collagen content appears to have decreased for some specimens and unclear in others. This is further supported by the fact that the failure load for collagenase treated specimens is significantly decreased compared to the control group, suggesting some collagen depletion is occuring. It must be noted that the current collagenase treatment was performed at room temperature rather than at body temperature (Dadgar et al., 1997) and this may have limited its potency. Alternatively, it is possible that collagen may still be present in the aortic wall even after being treated with collagenase (as seen histologically), but may not be coherent enough to be of any mechanical consequence. We know that human AAA have a significant amount of collagen remaining in the aortic wall despite decreased failure properties (Menashi et al., 1987; Rizzo et al., 1989), a possible explanation for the results we see here. If the collagen was not completely depleted in these tissues we would expect the tissue to be able to undergo a greater plastic strain than a tissue with no collagen because the latter would fail prematurely. Thus, the difficulty with collagen digestion may have overestimated the plastic strain, not underestimated it. A second limitation is the low resolution in displacement measurements owing to lack of an automated displacement data acquisition system during the time of these experiments. The accuracy of εpE- max may be affected slightly by the predicted pre-treatment length of the specimen since some specimen-to-specimen differences in clamping is likely to exist. A final limitation is that the specimens were not allowed to creep in the reverse direction once returned to its zero load length. Recent experiments show that a specimen returned from a creep cycle to zero load will continue to reduce in length while at that zero-load (reverse creeping). The plastic strain was typically measured within about 1 min of reaching zero load length. This will lead to a slight over estimation (about 3% according to recent preliminary studies) of the load-induced plastic strain. We do
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not believe these limitations will affect the overall finding since an overestimation would be a worst case scenario for the plastic strain. Thus the conclusion that the creation of an in vitro AAA via enzymatic treatments and mechanical loading is not feasible should remain valid. An in vitro AAA model can aid in many aspects of endovascular treatment of AAA. We found no evidence to support the feasibility of developing an in vitro AAA model by passively degrading the load bearing fibers of the aortic wall and subjecting these tissues to cyclic creep. Future attempts at developing an in vitro model may need to consider alternative approaches such as experimentation over longer durations (weeks and months, not days), in which the cells in the vascular wall are kept alive in a bioreactor environment and slowly exposed to proteolytic enzymes while subjected to gradually increasing cyclic pressure. Incorporation of tissue remodeling into an in vitro experimental setup is inevitable for development of in vitro AAA. We also found little evidence to support the postulate that passive creep may play an important role in aneurysmal dilation of a degraded aortic wall. REFERENCES
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