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1
Laboratory of Cardiovascular Physiology, Montpellier I University, Faculty of
Medicine, Av. Kennedy, Nîmes F 30907,
France.
2
Department of Thoracic and Cardiovascular Imaging, Montpellier University
Hospital, 34295-Montpellier Cedex, France.
Received July 12, 2000;
accepted after revision September 14, 2000.
Address correspondence to M. Dauzat.
Abstract
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SUBJECTS AND METHODS. Five self-expandable stents (Wallstent), five balloon-expandable noncovered Palmaz stents, and three balloon-expandable covered stents (Jostent) were placed in the infrarenal aorta of 13 New Zealand white rabbits. Systolic blood pressure changes, blood-flow velocity, systolic diameter, and diameter changes were measured and used to calculate the diameter compliance, the distensibility coefficient, and the pulsatility index.
RESULTS. Compliance (10-3 mm kPa-1) was 75.3 ± 20.1 before stenting and reached 94.7 ± 42.2 upstream, 38.8 ± 14.2 at the stent level (p < 0.05), and 70.8 ± 23.2 downstream from the stent. Distensibility (10-3 kPa-1) was 24.3 ± 6.3 before stenting and reached 27.8 ± 10.3 upstream, 10.5 ± 4.4 at the stent level (p < 0.001), and 21.9 ± 8.6 downstream from the stent.
Compliance and distensibility were significantly lower at the stent level than upstream and downstream (p < 0.05). Aortic diameter increased significantly at the stent level from 3.11 ± 0.40 mm before to 3.76 ± 0.42 mm after stenting. No significant difference was found among the three stent designs for all the studied data.
CONCLUSION. Regardless of the three tested stent designs, endovascular stenting produces a significant decrease in arterial wall compliance of the rabbit aorta.
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The rabbits received an IV injection of 200,000 U of penicillin G, 0.4 mg/kg of dexamethasone, and 500 U of heparin. The rabbits were intubated but breathed spontaneously and were placed in a supine position on a heating blanket. After subcutaneous injection of lidocaine (Xylocaine; Roger Bellon, Neuilly-sur-Seine, France), the rabbits underwent right femoral arteriotomy, and a 5-French introducer sheath with check-valve and side-arm (Radiofocus; Terumo, Tokyo, Japan) was introduced and placed in the abdominal aorta. The sidearm of the introducer was used to measure arterial blood pressure. The ECG and internal temperature were also monitored. Using a right transverse laparotomy, we exposed the infrarenal aorta, and the periaortic tissue was carefully dissected to place the sonographic transducers for diameter and blood flow measurement on the aorta. A 3.5 Short Magic Wallstent adapted to a mean vessel diameter of 3 mm was implanted in five rabbits, a balloon-expandable Palmaz stent with a 3- to 5-mm diameter and 7-mm length was implanted in five rabbits, and a balloon-expandable Jostent coronary stent graft with a 2.5- to 5-mm diameter range and 9-mm length was implanted in three rabbits. In all cases, the diameter of the stent and of the balloon catheter was adapted to the previously measured infrarenal aorta: a 3-mm diameter (n = 6) or a 3.5-mm diameter (n = 2). A Skinny dilatation catheter (Scimed, Maple Grove, MN) inflated at 6 atm was used to expand the eight balloon-expandable stents. Via the right femoral introducer, the delivery system of the Wallstent or the balloon-expandable stent mounted on the dilatation catheter was introduced over a 0.014-inch-diameter wire, and the stents were deployed in the infrarenal aorta with visual and biomicroscopic control. Repeated diameter and flow measurements were then obtained.
Diameter and Flow Measurements
Diameter and systolic diameter changes were measured in the 13 rabbits with
a 20-MHz emitting frequency 1-mm-diameter ultrasound echo tracking microprobe
(DMT201N; Crystal Biotech, Hopkinton, MA) mounted on a laboratory-made
silastic support connected by electric wire to an eight-channel sonometry
system (CBI 8000, Crystal Biotech) with a WT20 wall-tracker module and a
HVPD-20 pulsed Doppler module (Crystal Biotech). Spatial resolution of the
wall-tracker module, according to the manufacturer's specifications, is 1/8 of
the wavelength (i.e., 0.0077 mm at 20-MHz operating frequency). ECG, blood
pressure, diameter, and diameter changes were digitized with an MP 100 WSW
acquisition module (Biopac Systems, Goleta, CA) and processed with dedicated
software (AcqKnowledge version 3.5, Biopac Systems) on an IBM-compatible
computer.
Aortic blood-flow velocity was measured in 12 rabbits with a 20-MHz pulsed
Doppler probe. The mean arterial pressure was calculated as the time
averaged-mean arterial pressure. The pulsatility index was calculated as P/M
(P = total amplitude of the velocity curve, M = time averaged—mean
velocity). Diameter compliance was calculated as 2
d/P (P
= systolic — diastolic arterial blood pressure or systolic pressure
change) [10]. The
distensibility coefficient was defined as 2d/P/d. Measurements
were obtained before and after stent placement, 3 mm upstream, at the stent
level, and 3 mm downstream from the stent as shown in
Figure 1.
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Statistical Analysis
Descriptive statistics hereafter concern the pooled data of the 13 rabbits,
unless significant differences among groups were found. Aortic diameter,
compliance, distensibility coefficient, and pulsatility index change after
stenting, and differences among measurement sites (respectively upstream, at
the stent level, and downstream from the stent) were evaluated with the paired
t test. We compared the three stent-design groups by analysis of
variance with a Bonferroni test (Prism version 3.0; GraphPad, San Diego, CA)
(p values < 0.05 were considered significant).
Animal Care
The animal care complied with the "Principles of Laboratory Animal
Care," as formulated by the National Society for Medical Research, and
the "Guide for the Care and Use of Laboratory Animals"
[16].
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Diameter Compliance
In the entire rabbit population, diameter compliance (10-3 mm
kPa-1) was 75.3±20.1 before stenting and reached 94.7
± 42.2 upstream, 38.8 ± 14.2 at the stent level
(p<0.05), and 70.8 ± 23.2 downstream from the stent. At the
stent level, diameter compliance decreased from 77.63 ± 24.30 to 30.14
± 8.24 in the Wallstent group, from 69.10 ± 19.25 to 37.46
± 8.57 in the Palmaz group, and from 80.88 ± 18.61 to 55.25
± 17.83 in the Jostent group (p
0.01 for all groups). After
stenting, diameter compliance was significantly lower at the stent level than
upstream (98.48 ± 53.19 for the Wallstent, 85.36 ± 33.83 for the
Palmaz stent, and 103.13 ± 48.00 for the Jostent; p <
0.0005) or downstream (59.04 ± 13.83 for the Wallstent, 79.28 ±
30.80 for the Palmaz stent, and 75.53 ± 19.42 for the Jostent;
p < 0.0001). Diameter-compliance values measured upstream from the
stent for each stent design were not significantly greater than those obtained
before stenting. Diameter compliance values measured downstream from the stent
for the Wallstent and for the Jostent were not significantly smaller than
those obtained before stenting.
Distensibility Coefficient
In the entire rabbit population, the mean distensibility coefficient
(10-3 kPa-1) was 24.3 ± 6.3 before stenting and
reached 27.8 ± 10.3 upstream, 10.5 ± 4.4 at the stent level
(p < 0.001), and 21.9 ± 8.6 downstream from the stent.
There was a significant decrease in distensibility at the stent level after
stenting (p < 0.0001). The distensibility coefficient was
significantly lower after stenting at the stent level than upstream
(p < 0.0001) or downstream (p < 0.0001).
The distensibility coefficient upstream from the stent (27.47 ± 12.42 for the Wallstent, 28.49 ± 12.38 for the Palmaz stent, 27.59 ± 3.61 for the Jostent) for each stent design was not significantly greater than that before stenting (22.73 ± 5.02 for the Wallstent, 24.80 ± 8.49 for the Palmaz stent, 25.14 ± 6.99 for the Jostent). The distensibility coefficient downstream from the stent (15.92 ± 3.03 for the Wallstent, 24.86 ± 7.3 the Jostent) was not significantly lower than that before stenting.
Blood Pressure and Pulsatility Index
After stenting, the Doppler signal revealed no detectable change in
flow-velocity profile and no turbulence downstream from the stent. Pulsatility
index was 2.14 ± 0.54, mean arterial pressure was 7.89 ± 1.71
kPa, and systolic pressure change was 3.53 ± 0996 kPa before stenting
and showed no significant change at any site or among sites after
stenting.
Comparison of the three stent designs for the pulsatility index, diameter compliance, distensibility coefficient, aortic diameter, and diameter changes showed no significant difference before or after stenting at any site (Fig. 2A,2B,2C).
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Arterial Pulsatility
Stenting kept the arterial pulsatility index unchanged downstream and
upstream from the stented arterial segment and produced no downstream
turbulence, showing that stents did not significantly impede blood flow in
small-caliber arteries.
Compliance, Distensibility
Like the results of Rolland et al.
[14] for large arteries, we
found a compliance decrease in the stented segment of small-caliber arteries
in the rabbit aorta with the Wallstent and the Palmaz stent, whereas we found
no significant change upstream or downstream from the stented segment.
Therefore, independent of the three stent designs we tested, there was a
marked compliance mismatch between the stented and nonstented aorta. Unlike
the results reported by Rolland et al., no statistically significant diameter
compliance or distensibility coefficient difference among the three types of
tested stents was found before or after stenting at all the studied levels.
There are, indeed, marked differences between the study carried out by Rolland
et al. and our study. We Measured all parameters before and 15 min after
stenting in the same animal, without any significant change in hemodynamic
parameters during the entire procedure. Because we investigated much smaller
arteries than those of Rolland et al., we used a different technique with a
higher ultrasound frequency (20 instead of 10 MHz), ensuring a 0.01-mm spatial
resolution, and echo tracking instead of ultrasound time of flight, avoiding
any mechanical constraint on the artery, whereas Rolland et al. used a
silicone clip supporting two transducers positioned face to face around the
vessel. By using this different technique, we were able to show subtle changes
and differences in arterial wall mechanics in even smaller arteries (rabbit
iliac arteries) after laser-assisted microanastomosis compared with manual
microanastomosis [22]. On the
other hand, to avoid catheter-related flow disturbances, we did not perform in
situ intraarterial blood pressure measurements. Therefore, we could not
evaluate additional mechanical properties like hysteresis.
Aortic Diameter
The increase in aortic diameter after stenting is not the only factor
responsible for the decreased compliance at the stent level, as shown by the
significant decrease of the distensibility coefficient. The stent size was
chosen to fit the aortic diameter and to avoid overdilation. Some authors
purposely overdilated the stents by 30%, arguing that stretching of the
arterial wall is probably what usually occurs during stent placement for
occlusive vascular lesions in clinical conditions
[20]. In our study, we
mimicked the clinical practice with small-caliber arteries when the
nonstenotic artery upstream and downstream from the stenotic segment was
covered with the stent.
The analysis of diameter changes after stenting must consider structural and mechanical differences among stents. In experimental models and clinical trials, elastic recoil in vivo accounts for the loss of 10% of the initial post-stenting arterial diameter [23]. The Wallstent is an elastic stent in contrast with the plastic Palmaz and Jostent stents. Except for the recoil, the size of a stent correlates with the size of the balloon angioplasty catheter used for its deployment. Balloon angioplasty catheters designed for the deployment of expandable stents are not compliant and support high inflation pressures. This feature accounts for a correct expansion of the balloon despite the parietal strain usually encountered in clinical practice. The stent diameter will not increase unless additional angioplasty with a larger balloon diameter is performed. In contrast, because of the elastic properties of the Wallstent, there is a spectrum of possible sizes for a given diameter. With the Wallstent, the final diameter depends on the strain applied by the arterial wall itself and surrounding tissues.
The possible decrease in parietal strain due to anesthesia and periaortic dissection could explain the fact that we found a larger (although not statistically significant) aortic diameter after stenting with the Wallstent than with the Palmaz stent or the Jostent. Balloon-expandable stents appear to offer the advantage of an expansion precisely adjusted to the limits of the delivery balloon.
Covered Stents
A higher rate of both acute thrombosis and late restenosis has been
reported when placing covered stents instead of noncovered stents in the
rabbit aorta [24]. For Tepe et
al. [24], stent failure could
result from a decreased biocompatibility of the coverage. Covered stents seem
to exhibit the same mechanical properties as uncovered stents, but their
number in our study was too low to draw any definitive conclusion.
Limitations of Our Study
The differences in stent design concern the type and size of the wires, the
number of strut—strut intersections, the length and surface of the
stent, and its radial strength and longitudinal flexibility
[23,
25]. Moreover, some parameters
cannot be studied independently, so it is not possible to incriminate one
single parameter for a given observed difference. However, our study reflects
the clinical practice in which shorter Wallstents are not available and long
Palmaz stents are too rigid to reach some lesions in tortuous vessels. Such
differences related to stent design concern all comparative studies performed
with the stents currently used in clinical practice
[9]. In a 1999 publication,
researchers established the stented segment length as an independent predictor
of restenosis [26]. In our
study, the Wallstent stents we placed in the aorta were nearly twice as long
as the two other types of stents. This difference may be reflected by the
insignificant decrease in compliance and pulsatility in our study.
Nevertheless, the length of the stented segment was always less than 2 cm in
our study so that all our cases corresponded to the same group in the
classification used in the study of Kobayashi et al.
[26].
Our study did not assess the histologic changes, including atrophy of the media and the neointima formation, that occur from 1 day to 4 weeks after stent placement, whereas the maximum restenosis rate occurs 8 weeks after stent placement and remains unchanged thereafter [18]. Although chronic studies would be interesting, stenting is the first causal factor, and its immediate mechanical consequences are thought to be the initial cause of histologic changes that may, in turn, induce or worsen mechanical changes [6].
The invasive procedure we used for the placement of ultrasound probes represents another limitation of our study and would not be suitable for chronic studies. As previously reported, both the anesthesia and the surgical dissection of the aorta modify wall mechanics [27, 28]. We are now developing and validating a new method for chronic studies using transparietal B mode sonography and a dedicated image-analysis software to noninvasively evaluate the diameter changes.
Finally, the rabbit aorta is an elastic artery, and the results of our study cannot be directly extrapolated to the treatment of small-caliber arteries in humans. Nevertheless, similar histologic reactions to stenting have been reported by Robinson et al. [29] in both the rabbit aorta and the pig or dog coronary artery, and the rabbit aorta is widely used as a model for the study of stent-induced histologic changes.
To our knowledge, ours is the first in vivo study comparing the effects of different stent designs on wall mechanics of small-caliber arteries. We found no significant differences among the Wallstent, the Palmaz, and the Jostent. At implantation, they reduce the compliance of the stented arterial segment and induce a compliance mismatch. As previously suggested, this compliance mismatch may be one of several factors promoting restenosis. However, long-term studies are mandatory for a thorough understanding of the underlying mechanisms.
Acknowledgments
We thank Margaret Manson for her help in revising the English
manuscript.
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