© 1997 American Heart Association, Inc.
Articles |
In Vivo/In Vitro Comparison of Rat Abdominal Aorta Wall Viscosity
Influence of Endothelial Function
From the Départements de Pharmacologie (P.B., S.L.) and Médicine Interne (M.S.), Hôpital Broussais, and INSERM U337 (P.B., Y.B., P.L., P.C.-C., A.B., J.F.R. de la F., M.S., S.L.), Paris; and URA-CNRS 879, Saint Cyr l’Ecole (P.C.), France.
Abstract
Abstract Arterial wall viscosity (AWV) is a potential source of energy dissipation in circulation. That arteries, which are known to be markedly viscous in vitro, have lower viscosity in vivo has been suggested but not demonstrated under similar pressure conditions. Endothelium, which may modulate AWV through smooth muscle tone, could contribute to the low level of viscosity in vivo. Our objectives were first to compare AWV of the rat abdominal aorta, in vivo and in vitro, with similar pulse-pressure waves, and second, to determine whether endothelial function influences AWV in vivo and in vitro. The diameter of the abdominal aorta and distending pressure were measured in vivo and in vitro with a high-resolution echotracking system and a micromanometer, respectively. AWV was calculated as the area of the pressure-volume curve hysteresis. After in vivo examination, the arterial segments were isolated in vitro and submitted to resynthesized pressure waves identical to those recorded in vivo. Deendothelialization was performed in vivo by balloon rubbing; then arteries were examined either in vivo or in vitro. AWV was markedly lower in vivo than in vitro (6.6±0.7 versus 22.7±3.7 J · m-1 · 10-5, respectively; P<.001). After deendothelialization, a sustained 40% increased AWV was observed during a 15-minute follow-up (P<.01). In vitro, deendothelialized arteries have a 64% higher AWV than segments with endothelium (P<.01). Our results indicate that the physiological effective viscosity, measured in vivo in intact animals, is threefold lower than the intrinsic viscosity of the arterial wall, measured in vitro. Endothelium removal determines a sustained increase in AWV, either in vivo or in vitro. These results suggest that active mechanisms compensate for intrinsic viscosity under physiological conditions. One of these energy-saving mechanisms might be dependent on normal endothelial function.
Key Words: viscosity • endothelium • arterial compliance • aorta
Viscoelasticity has long fascinated physiologists, because of the belief that the underlying mechanisms are the key to the fundamental mechanisms for generating smooth muscle force.1 Although the viscoelastic behavior of large arteries has been extensively investigated, few studies have focused on viscosity itself. The ejection of the heart is pulsatile, while peripheral circulation is continuous. Large artery function is not only to conduct the blood from the heart to the periphery but also to store the potential energy generated by the heart during systole2 3 4 and restore it during diastole. Thus, large arteries can preserve the potential energy (ie, pressure) from attenuation along the arterial tree so that driving pressure remains high enough at the periphery to drive the blood in resistive networks. During systole, the elastic behavior of the arterial wall allows the internal diameter to increase proportionally to blood pressure. The stored energy is proportional to the product of strain and stress.5 Purely elastic materials allow the whole stored energy to be restored during the unloading phase. However, biological tissues in general, and arteries in particular, are not purely elastic but exhibit a markedly viscous behavior. Vascular tissues exhibit greater dimension at any given pressure during unloading than during loading phase, which determines a hysteresis loop on the pressure-volume relationship. The area of the hysteresis loop has a dimension of energy. Thus, while part of the energy stored by the arterial wall during elastic distension is fully restored, the remaining part of the energy corresponding to the viscous deformation (area of the hysteresis loop) is dissipated within the arterial wall.5 6 7 Thus, AWV is considered as the main source of pressure-wave attenuation along the arterial tree in theoretical models.8 In these respects, AWV, an energy-dissipating mechanism, can reduce the efficiency of heart-vessel coupling.5
AWV has mainly been determined in vitro1 9 10 and is easily measurable. The determinants of AWV are the amplitude and frequency of pulsatile stress, together with arterial smooth muscle content and its degree of contraction.2 11 12 AWV was very small in the few published in vivo studies,4 6 7 13 14 despite the presence of physiological stimuli reported to increase AWV. It has not been clearly determined whether AWV is smaller in vivo than in vitro at the site of conduit arteries. Furthermore, AWV may have been underestimated in vitro. Indeed, as viscosity depends on the frequency and velocity of stretching,2 the low-amplitude sinusoidal pressure waves used in in vitro experiments may not affect arterial wall behavior in the same way as the complex pressure waves observed in vivo.1 To our knowledge, AWV values have not been compared in vivo and in vitro with similar pulse-pressure waves.
Various interdependent factors may contribute to the apparent in vivo/in vitro difference in AWV, including the pulse-pressure wave pattern, innervation and vascularization of the arterial wall, pulsatile flow and shear stress, and smooth muscle tone. In particular, viscosity correlates with the activity of smooth muscle in a given artery.1 2 This activity results from complex interactions between the endothelium, smooth muscle cells, and humoral and mechanical factors.15 16 The release of endothelium-derived relaxing factors is modulated by the amplitude and frequency of flow, pulse pressure,17 18 and mechanical deformation,19 while the endothelium modulates stretch-induced sustained and phasic myogenic responses.16 20 21 22 In this study, we focused on the influence of the endothelium on AWV.
We first compared AWV at the site of a large muscular artery (the rat abdominal aorta), in vivo and in vitro, with similar pulse-pressure waves, and then investigated the potential influence of endothelial function on AWV in vivo and in vitro.
Methods
Animals
We used 27 male Wistar rats (Iffa-Credo, France) aged 12
weeks and weighing 320±20 g (mean±SD). The animals were managed
according to French Ministry of Agriculture guidelines. Experiments
were conducted under disodium thiopental anesthesia (50
mg/kg IP). Animals were killed by exsanguination.
The comparison between in vivo and in vitro viscoelastic properties of
rat abdominal aorta was performed in six rats during a first
experiment, in which AWV (Fig 1
) was first determined in vivo and then
in vitro on the same arterial segments. The role of
endothelium was studied, under in vivo and in vitro
conditions, in a second and third experiment. In the second experiment,
nine rats were studied in vivo, before and after abdominal aorta DE,
and compared with five sham rats. In the third experiment, DE of the
abdominal aorta was performed in vivo in seven animals; the aortas were
then studied under in vitro conditions and results compared with those
of in vitro controls with intact endothelium, matched
for mean and pulse pressures.
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Experiment 1: In Vivo and In Vitro Determinations of AWV
In Vivo Determination
The 2F microtransducer was introduced through the left femoral
artery of anesthetized animals for aortic blood pressure
measurements. A median laparotomy was then performed and the AA was
exposed but not dissected. A 10-MHz focalized transducer was
stereotactically positioned 1 cm above the aortic
trifurcation, using warm (37°C) isotonic saline as coupling medium.
The position of the pressure probe was adjusted visually and assumed to
be correct when the tip of the catheter was detected on the
radiofrequency signal of the aorta. Pulsatile changes in diameter and
pressure were then measured during three periods of 4 seconds each.
In Vitro Determination
The experimental setup developed for in vitro determination of
AWV is represented in Figs 2 through 4. Arterial diameter and
pressure were measured and computed in the same way as in the in vivo
determination.
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Abdominal aortic segments were collected immediately after in vivo
investigations. The collateral side branches were ligated carefully
under a binocular microscope and the AA was gently separated from the
inferior vena cava. Two 3-0 silk laces were then placed
around the AA; the first at the level of the left renal artery, and the
other at the level of the aortic termination. The length of the AA was
measured in situ between the laces with a two-point compass. The
segment of artery was quickly removed and placed in ice-cold Krebs
buffer (118 mol/L · 10-3 NaCl,
4.6 mol/L · 10-3 KCl, 2.5
mol/L · 10-3 CaCl2
2H2O, 1.16 mol/L ·
10-3 MgSO4, 1.16 mol/L
· 10-3 KH2PO4, 11.1
mol/L · 10-3 glucose, 25
mol/L · 10-3 NaHCO3,
95% O2, 5% CO2). The explant length was
measured and shrinkage (35±5%) calculated. The remaining side
branches were ligated. The proximal end of the AA was cannulated with a
2F cannula and connected to a perfusion line. The distal end was
cannulated with a 2F high-fidelity microtransducer to measure
intraluminal pressure. The arterial segment was then
mounted in a specially designed organ chamber containing
oxygenated Krebs buffer at 37°C. The AA segment was then
pressurized with Krebs, extended to its in vivo length, submitted to a
constant pressure of 100 mm Hg, and allowed to equilibrate for 30
minutes. The leakage was checked by temporarily emptying the chamber
and estimating the speed of refilling. When leakage exceeded 1 mL/min
and could not be corrected, the experiment was discarded. Pressure was
then set at the mean arterial pressure in vivo for 15
minutes, and the artery was connected to the pressure-wave synthesizer
(described above). The perfusion line was composed of low-compliance
polyethylene tubing, a pressure reservoir containing
oxygenated Krebs buffer at 37°C, and a proximal
adjustable Windkessel segment. A rectangular negative ramp signal was
found to provide the best fit for the morphology of the pressure wave
observed during the corresponding in vivo experiment (Fig 5). The pressure-signal morphology was
fine tuned by careful adjustment of the Windkessel chamber size.
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The amplitude and frequency of pulse pressure were adjusted to the pulse pressure measured in vivo for each rat. The artery was submitted to this pulsatile pressure regimen for 5 minutes, and pressure and diameter were then measured for 4 seconds. To ensure that the frequency content of the distending pressure in vitro was identical to that in vivo, we performed a Fourier analysis (8192 samples) of the pressure waves under both conditions and compared them afterward.
Experiment 2: Effects of DE on AWV In Vivo
DE was performed by means of the balloon catheter procedure
originally described by Clowes et al.23 As our objective
was to determine the maximal DE not altering arterial wall
function, we slightly modified the classical method to avoid
overstretching the aortic wall and checked that aortic ring contraction
and relaxation were preserved (see below).
The animals were prepared as described in experiment 1, and the pressure-diameter relationship was measured at baseline. A 2F Forgarty catheter (Baxter) was then introduced through the left common carotid artery and gently pushed down through the thoracic aorta toward the abdominal aorta. When the tip of the catheter was located near the area of ultrasonic diameter measurement, it was pushed 1 cm further. The balloon was then inflated with the minimal volume of air (usually 50 µL) required to obtain occlusion of the artery (confirmed by the disappearance of the downstream pressure signal). The catheter was then slowly withdrawn until resistance was felt at the level of the diaphragm. The procedure was repeated once, and the catheter was then deflated and withdrawn. In sham animals, the catheter was pushed until it reached the site of diameter measurement but was not inflated. The pressure-diameter relationship was measured 1, 5, and 15 minutes after the procedure.
To ensure that the frequency content of the blood pressure after DE was identical to that before DE, we performed a Fourier analysis (8192 samples) of the pressure waves under both conditions and compared them afterward.
To check that our DE procedure damaged the endothelium without altering smooth muscle function, aortic rings were studied in isometric conditions. The abdominal aorta was removed from five DE and five sham animals and placed in ice-cold Krebs buffer. The arteries were then dissected and cut into rings 3 to 4 mm long. Two rings from each animal were placed in an isometric organ chamber (Marty Technology) and allowed to equilibrate at 37°C in oxygenated Krebs buffer. Segments were exposed to 60 mol/L · 10-3 KCl for 15 minutes and then to cumulative concentrations of norepinephrine (10-9 to 10-5 mol/L). When the maximal contractile response was obtained, endothelium-independent relaxation was tested by adding 10-5 mol/L papaverine. The maximal force developed under KCl stimulation and the degree of endothelium-independent relaxation were taken as indices of smooth muscle function. The aortic rings were then contracted with 3x10-7 mol/L norepinephrine and exposed to cumulative doses of carbamylcholine (10-9 to 10-5 mol/L). The capacity of aortic segments to relax in a carbamylcholine concentration–dependent manner was used as an index of endothelial integrity.15 Data for each concentration are expressed as the percentage of maximum baseline contraction.
Experiment 3: Effects of DE on AWV In Vitro
DE was performed in vivo by means of a balloon catheter,
following the procedure described in experiment 2. Then, abdominal
aorta segments were collected in the same manner as in experiment 1 and
studied in vitro as described above. The mean and pulse pressures
applied in vitro to these arterial segments were chosen
identical to those measured in vivo previously from the DE
procedure.
Determination of the Dynamic Diameter-Pressure
Relationship
As viscosity is measured by the area of the hysteresis of the
diameter-pressure relationship (Fig 1
), we first
simultaneously measured pulsatile changes in diameter and
pressure to determine their relationship. In vivo and in vitro
measurements were made in the same way. Pulsatile changes in diameter
were measured with the NIUS 1 echotracking device (described below).
Pulsatile changes in distending pressure were measured using a 2F
high-fidelity microtransducer (Millar Instruments, Nycomed; Fig 4).
Aortic diameter was measured by the NIUS 1 echotracking device (ASULAB Research Laboratory), which has been fully described elsewhere.24 25 26 Briefly, the device measures internal diameter and its systolic-diastolic variations with a precision close to 50 µm and 1 µm, respectively. This high precision is made possible by the quadratic fit of a specific window of the radiofrequency signal, the high frequency (75 MHz) of signal tracking, and the digitization of this signal on 8 bits. The software was slightly modified for viscosity measurements. Diameter and pressure were simultaneously digitized and stored at a frequency of 2.5 kHz, giving a 4x10-4-second time domain definition (0.0125 radian for a frequency of 5 Hz).
The Doppler technique was used to place the probe perpendicularly
to the arterial axis, in its largest cross-sectional
dimension. After the transducer was switched to radiofrequency mode,
the backscattered echoes from both the anterior and posterior walls
were visualized on an oscilloscope. The radiofrequency signals of both
walls exhibiting a high signal-to-noise ratio were easily tagged by two
electronic trackers, so that their movements could be monitored
continuously. Recordings of the tracker movements were used to
derive a digitized signal of arterial diameter variations.
Distending pressure was measured simultaneously at the same
position. The computerized acquisition system derived the lumen
cross-sectional area-pressure curve from the two continuous signals of
arterial diameter and pressure and fitted it using an
arctangent model with three independent parameters (
,
ß, and 
), as described by Langewouters et al27 :
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, ß, and have a physiological
meaning.27 represents the maximal LCSA divided
by 
/2, ß is the pressure at which the compliance is maximum, and
is the pressure at which compliance is half the maximum.
Various definitions of compliance are used in the
literature.28 In the case of a cylindrical vessel,
arterial cross-sectional compliance (C) is defined by the
change in cross-sectional area (
LCSA) for a given change in
intravascular pressure (P). Owing to the nonlinearity of the
cross-sectional area–pressure curve, compliance falls as pressure
increases. It is therefore important to establish the
compliance-pressure curve in the systolic-diastolic
range (dynamic cross-sectional compliance). This value was obtained by
deriving equation 1 with respect to pressure as follows:
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refers to dynamic cross-sectional compliance.
Arterial cross-sectional distensibility (Dist) is defined
by the relative change in cross-sectional area (LCSA/LCSA) for a
given change in distending pressure and represents compliance
normalized to the corresponding cross-section of the artery:
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Evaluation of AWV
AWV can be estimated by the hysteresis of the pressure-volume
relationship or by the ratio of the dynamic-to-static elastic
moduli.5 12 We focused on the hysteresis loop of the
pressure-LCSA relationship, as the ratio of the dynamic-to-static
elastic moduli cannot be easily determined in vivo for obvious
reasons.
The energy per unit of arterial length exchanged from the
blood to the arterial wall during distention can be written
as:
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L is an arbitrary length, P is
pressure, and dLCSA is the elementary variation in lumen area.
AWV determines the emergence of a hysteresis loop (Fig 1), in which it is possible to
distinguish the "distention" limb (A
B) during systole from the
"retraction" limb during diastole (BA).
Arterial diameter is smaller during distention than
retraction at a given distending force. The hysteresis loop corresponds
to a classical adiabatic cycle in thermodynamics. Indeed, the energy
stored during the distention phase (WE) is not fully restored during
retraction. This energy dissipated by the arterial wall
(dW) per unit length during one cycle (dL) represents Wv and
can be written as:
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corresponds graphically to the
area of the hysteresis loop of the pressure-LCSA relationship (Fig 1).
We developed dedicated software to measure this area directly from the
original pressure and diameter recordings, using an integrative
algorithm. Energies are expressed in joules per meter during one cycle.
Wv is expressed either in absolute values or as a percentage of WT,
with WT=Wv+WE. This software permitted us to detect a phase lag as
small as 10-5 radian on simulated sinusoidal
time series. The practical precision of this method is given by the
precision of pressure and diameter measurements (0.5 mm Hg and
1 µm, respectively) and can be estimated as
>1x10-10 J ·
m-1. Viscosity was measured on all cardiac
cycles and was averaged for one cycle.
In Vitro Validation of Viscosity Measurements
A crucial point in the measurement of the hysteresis loop
of the pressure-LCSA relationship is the synchronization of the
pressure and diameter signals. We therefore developed an experimental
setup (Fig 2) to estimate the potential
time delay between these two signals. A two-entrance low-compliance
chamber was closed with a purely elastic plastic membrane and filled
with distilled water. A high-fidelity microtransducer was positioned
inside the chamber immediately beneath the membrane, and its motion was
measured with the echotracking NIUS system. The whole system was
submitted to various pressure waveforms generated by a homemade
pressure wave synthesizer. This device was composed of three elements.
A function generator (Hewlett Packard) was programmed to generate
appropriate electrical signals. These signals were amplified by a solid
state power amplifier (Classé Audio 70) and fed into a powerful
low-frequency loudspeaker (Focal) coupled mechanically to a
low-friction microsyringe (Hamilton). This pressure synthesizer could
generate pressure signals adjustable in amplitude, frequency, and
shape. We submitted this elastic membrane to various pressure waveforms
(square, sinusoidal) of various frequencies (1 to 50 Hz). Any
desynchronization in the acquisition or processing of the pressure and
diameter signals was detected by the appearance of a hysteresis loop on
the pressure-displacement curve. Both transducers were found to be
linear in the frequency and amplitude range observed in animals. The
two signals were simultaneously acquired, and no time delay
was observed. As a result, the hysteresis loop area was negligible
(<10-7 J/m). To illustrate the
synchronization and linearity of our system, a complex waveform
associating a carrying wave of low frequency and high amplitude (2 Hz,
40 mm Hg, respectively), modulated by a high frequency and small
amplitude (30 Hz, 10 mm Hg, respectively). The resulting pressure
and displacement waves and the pressure-displacement relationship are
represented in Fig 3. Despite
the wide range of frequency and amplitude present in the pressure
signal, no hysteresis was discernible.
Statistical Analysis
Data are expressed as mean±SEM. In vitro measurements were
compared with in vivo measurements by using a paired
nonparametric Wilcoxon’s test. The effects of DE
on arterial wall elasticity and viscosity were
analyzed by using repeated-measures ANOVA, with time as the
within-subject factor and the group (DE, sham) as the between-subject
factor. The C matrix post hoc test was used to test changes in the
parameters versus baseline values. All tests were bilateral
and values of P<.05 were considered to denote significant
differences. All statistical tests were done on SYSTAT
software.29
Results
Experiment 1: In Vivo/In Vitro Comparison of Viscoelastic
Properties
The frequency content of the in vitro pressure waves was not
significantly different from that of in vivo waves (harmonic 1:
6.12±0.45 versus 5.54±0.49 mm Hg; harmonic 2: 3.60±0.32 versus
2.89±0.32 mm Hg; harmonic 3: 1.54±0.20 versus 1.42±0.25
mm Hg; and harmonic 4: 0.66±0.15 versus 0.69±0.10 mm Hg).
LCSA was lower in vivo than in vitro, while distensibility and
compliance were markedly higher in vivo (Table 1). AWV was markedly lower in vivo than
in vitro, both when expressed in absolute values and when normalized to
the total work exchanged. Typical pressure-LCSA relationships obtained
in vivo and in vitro in the same artery are represented in
Fig 6. The frequency content of the
distending pressure in vitro did not differ from that observed in vivo,
since the first four Fourier components were not significantly
different between conditions.
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Experiment 2: Effect of DE on Arterial
Parameters In Vivo
Arterial parameters did not differ between
sham and DE animals at baseline (Table 2). In sham rats, as no significant
time-dependent change in arterial parameters
was observed, the four (baseline, 1, 5, and 15 minutes) LCSA-pressure
and distensibility-pressure curves were pooled (Figs 7 and 8).
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A significant decrease in systolic, mean, and
diastolic blood pressures was observed in DE animals 1
minute after DE, but not at later times. No change in pulse pressure or
heart rate was observed (Table 2). The frequency content of the
distending pressure after DE did not differ from that observed before
DE, since the first four Fourier components were not significantly
different between conditions during follow-up. The pressure-diameter
curve of DE animals was shifted upward by the procedure (Fig 7). At a
blood pressure of 125 mm Hg, the increase in aortic LCSA was
significant 1 and 5 minutes after DE, but not at 15 minutes.
Distensibility was marginally altered by DE, and the
distensibility-pressure curve was slightly shifted downward: at a blood
pressure of 125 mm Hg, there was a slight but significant
decrease in distensibility 1 minute after DE. Aortic compliance,
calculated at 125 mm Hg, was not modified 1 minute after DE, but
was significantly increased thereafter.
The total work exchanged during the cardiac cycle (WT) did not change
during the procedure in DE or sham rats. An important observation was
that DE induced a significant increase in AWV (Wv, +41%; WP, +43%),
which occurred at the first minute and persisted thereafter (Table 2).
A representative pressure-LCSA relationship trace
obtained before and after DE is shown in Fig 9.
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Experiment 3: Effect of DE on Arterial
Parameters In Vitro
DE arteries showed lower diameter, distensibility, and compliance
than controls with intact endothelium when examined in
vitro (Table 3). The markedly reduced
distensibility of DE arteries was accompanied by a decreased ability to
store energy during distention: the absolute energy dissipated in
viscosity was smaller than in controls. In addition, the percentage of
exchanged energy dissipated in viscosity was 64% higher than in
controls.
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Endothelial and Smooth Muscle Cell
Function
Balloon rubbing of the abdominal aorta led to complete inhibition
of carbamylcholine-induced endothelium-dependent
relaxation (Fig 10). This inhibition
was obtained with no detectable alteration of maximal
endothelium-independent contraction or relaxation.
Indeed, maximal contraction in response to 60 mol/L ·
10-3 KCl did not differ between DE and sham
animals (2.94±0.23 g versus 3.17±0.18 g, respectively). In addition,
endothelium-independent relaxation induced by
papaverine was not significantly different in DE and sham animals
(-124±6% versus -106±4%, respectively).
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Discussion
Large arteries not only conduct blood from the heart to the periphery but also store energy during systole, restore it during diastole, and transmit it along the arterial tree. This ability to store and transmit energy, which is the main determinant of cardiovascular efficiency, is directly influenced by AWV.
The two main findings in this study are that (1) AWV is lower in vivo than in vitro under similar pulsatile conditions and (2) DE leads to a marked increase in AWV either in vivo or in vitro.
In Vivo/In Vitro Comparison of Viscoelastic Properties
Consideration of Methods
Two prerequisites for this study were accurate measurement of AWV
in situ, under physiological pressure and flow
conditions, and the comparability of these measurements with those
obtained on the same arterial segments in vitro. We had to
make theoretical choices as to the characterization of AWV. Indeed,
although many models of viscoelasticity have been proposed, none has
proved fully able to describe the behavior of the vascular
wall.30 Thus, instead of using a model, we took advantage
of the ability of the NIUS system to accurately determine the
pressure-diameter relationship and chose to measure AWV directly,
through its thermodynamic expression, ie, the hysteresis loop area of
the pressure-diameter curve. This approach has the advantage of
stressing on the physical consequence of viscosity, which is to
dissipate energy. This method was initially applied by
Brodley6 and Bertram7 to arterial
wall mechanics. As discussed in "Methods," a prerequisite for
measuring the hysteresis loop of the pressure-diameter relationship was
synchronization of the two signals. In a dedicated experimental setup
including a highly elastic membrane, we showed that this was
effectively the case, with a negligible hysteresis loop area.
The second challenge was to generate pressure waveforms in vitro as
close as possible to those recorded in vivo. Previous works on AWV
have involved low-amplitude sinusoidal pressure waveforms because of
fundamental and technical problems. In these conditions of
low-amplitude deformation, AWV is fairly insensitive to
frequency,30 although the amplitude of stress
oscillations was not controlled.3 Moreover,
viscous phenomena depend not only on pulse-pressure frequency but also
on its amplitude and velocity of change.12 The natural
pressure waveform exhibits both high-amplitude and high-frequency
components, which predominate during the distention phase. We thus
developed a pressure wave synthesizer to generate a pressure waveform
similar to the natural counterpart. A rectangular negative ramp signal
provided the best fit to the morphology of the pressure wave observed
during the corresponding in vivo experiment (Fig 5). The pressure
signal morphology was fine tuned by careful adjustment of the
Windkessel chamber size. The amplitude and frequency of pulse pressure
were adjusted to the pulse pressure measured in vivo for each rat.
Because of the time constraints, this adjustment was done by "visual
fit" during the experiment and checked by harmonic decomposition
afterward. As presented in the "Results" section, the
first four Fourier components of the pressure waves were not
significantly different between in vivo and in vitro baseline
conditions, and no correlation between the harmonic content of pressure
and viscous parameters was found.
Consideration of Findings
One of the main findings is that AWV was lower in vivo than in
vitro. Such a difference had only been suspected, on the basis of
comparisons of experiments on different species, at different
arterial sites reporting in vivo or in vitro
data.3 5 7 9 10 13 In contrast, we observed this in
vivo/in vitro difference in AWV in the same animals and at the same
arterial site.
Although the experimentally generated pulsatile waveforms were similar to their natural counterparts, flow and smooth muscle tone were absent. The lack of smooth muscle tone may explain the larger in vitro aortic diameter but is unlikely to have led to an overestimation of AWV, as smooth muscle contraction is generally reported to increase AWV.1 2 12 The impact of the absence of flow on viscosity in vitro is more difficult to estimate. Arterial blood flow and smooth muscle tone may have opposite influences on AWV, the former decreasing it through flow-dependent vasorelaxation and the latter increasing it.2 Other mechanisms potentially explaining the in vivo/in vitro difference in AWV include an interruption of blood flow to the vasa vasorum,31 leading to ischemic changes. Although our experiments were conducted in oxygenated medium, we could not rule out transient ischemia and subsequent functional changes. The replacement of blood by Krebs buffer may also have had an influence. In addition, diffusion of electrolytes, water, and macromolecules into the arterial wall may have changed its viscoelastic properties.32 In contrast, an alteration in endothelial function in vitro can be ruled out, as the segments of abdominal aorta submitted to pulsatile pressure were prepared in the same way as in the isometric experiments (which indicated normal relaxation in response to carbamylcholine).
Whatever the mechanisms underlying the in vivo/in vitro difference in AWV, the most important point is that AWV is very low in vivo. AWV is generally described as a passive phenomenon due to the physical properties of arterial wall components.30 This passive phenomenon was best expressed in our in vitro conditions, as the arterial segment was isolated from neurohumoral and flow-dependent control. The fact that AWV is very low in vivo suggests that active mechanisms compensate for intrinsic viscosity in physiological conditions. Such active mechanisms would imply dynamic behavior of arterial wall components during the cardiac cycle, minimizing AWV. As viscosity dissipates energy, these mechanisms (including preserved endothelial function) may improve the efficiency of the cardiovascular coupling energy balance.
Arterial Parameters After DE
Smooth muscle tone results from complex interactions between the
endothelium, smooth muscle cells, and mechanical
factors. In particular, endothelium-derived relaxing
factor release is modulated by the amplitude and frequency of flow,
pulse pressure, and mechanical deformation. On the other hand,
endothelium modulates stretch-induced sustained and
phasic myogenic responses.16 19 20 22 This suggests that
endothelial function could interfere with the dynamic
behavior of the vessel wall and specifically with AWV.
Consideration of Methods
We used the classical balloon method to suppress the
endothelial layer locally at the site of
pressure-diameter measurements. We took special care to minimize any
damage to the arterial wall caused by balloon
overinflation, and no visible arterial dilation occurred.
However, a degree of smooth muscle dysfunction cannot be ruled out,
even if gross contraction and relaxation functions were preserved when
studied isometrically in vitro.
DE may modify AWV through changes in the pressure waveform. Indeed, during the denudation procedure, the aortic pressure proximal to the occlusion likely increased and that distal to the occlusion decreased. These changes in hemodynamic conditions could alter the aortic pressure waveform, even during the follow-up period. To ensure that the frequency content of the distending pressure after DE was identical to that before, we performed a Fourier analysis of the pressure waves under both conditions. We checked that the first four Fourier components of the pressure waves were not significantly different between baseline and post-DE, under in vivo conditions. Therefore, the time-dependent change in AWV after the denudation may well be a true change in viscosity and not a reflection of changes in the waveform of aortic pressure. DE arteries, studied in vitro, were submitted to a pulsatile pressure regimen, the amplitude and frequency of which were adjusted to the pulse pressure measured in vivo after DE for each rat.
Consideration of Findings
The main finding in the DE experiments is that DE causes a large
and sustained increase in AWV in vivo and that this increase in AWV is
further increased when the DE arteries are studied in vitro. In vivo,
the hemodynamic modifications that accompanied the DE
procedure were transient and minimal and were thus unlikely to explain
the increase in AWV, by contrast with the work of Pagani et
al,14 in which the increased hysteresis after nitrix oxide
donor administration in sheep was accompanied by large
hemodynamic changes. Other mechanisms potentially
explaining this increase in AWV include a rise in arterial
smooth muscle tone after DE, but this is very difficult to assess in
situ.
The relationship between changes in LCSA and distensibility after DE is less obvious for in vitro experiments than for in vivo ones. For instance, transient and opposite changes in LCSA and distensibility occurred in vivo after DE, the former increasing and the latter decreasing. The influence of blood pressure was minimized by the expression of distensibility as an isobaric parameter (Dist125). It is well accepted that a higher LCSA redistributes the loads from the more distensible to the less distensible components of the arterial wall, resulting in a lower distensibility. In vitro, the relationship between the low LCSA and the low distensibility is more difficult to explain, since a low LCSA is supposed to redistribute the loads from the less distensible to the more distensible components of the arterial wall, resulting in a higher distensibility, in a classical model of the arterial wall in which the elastin and collagen fibers are in parallel.33 A more complex model, suggested by O’Rourke et al,34 is that at normal distending pressure, smooth muscle in the arterial wall is in series with some of the stiffer collagenous components but in parallel with the elastic lamellae. Contraction of smooth muscle tenses the collagenous components, thus decreasing distensibility, whereas dilation transfers stresses to the elastic lamellae, thus increasing distensibility. Other mechanisms may be involved.
The high level of viscosity of in vitro DE arteries is accompanied by a very low distensibility compared with in vitro controls. As stated above, an increased AWV may not only increase hysteresis but also reduce distensibility. By nature, a viscous material is stiffer under dynamic pressure regimen, while it dissipates more energy (ie, through increased hysteresis).
In the present work, AWV was increased in vitro, with and without endothelium, while changes in lumen size were opposite. It is therefore unlikely that the increase in viscosity relates only to change in lumen size, shifting the load bearing from distensible to stiffer components of the arterial wall. In addition, although changes in LCSA and distensibility may result from changes in arterial smooth muscle tone, changes in AWV may result from other factors. Indeed, in a preliminary in vitro experiment using the same methodology, AWV was not significantly modified by either phenylephrine or sodium nitroprusside in arterial segments having functional endothelium, despite significant vasoconstriction and vasodilation, respectively.
Two other mechanisms besides smooth muscle contraction may possibly contribute to the increase in AWV after DE: overstretching in response to an abrupt and durable rise in distending pressure, which induces a rearrangement of wall components35 and arterial inflation and suppression of the endothelial barrier, which may promote electrolyte, water, and macromolecule diffusion into the arterial wall. On the contrary, endothelium could actively promote a fast adaptation of the vessel wall to rapid changes in stress (within the cardiac cycle), possibly through the release of NO in response to pulsatile pressure and/or pulsatile flow.
Although the precise mechanism underlying the increase in AWV after DE is unclear, our results indicate that an intact endothelial layer is necessary to keep AWV low. However, despite the large increase after DE, AWV remained lower than in vitro. This finding suggests that factors other than the endothelium are involved in the in vivo/in vitro AWV difference. In addition, it suggests that the endothelial layer, although functional in vitro, cannot bring AWV down to its in vivo level.
The importance of energy dissipation through viscosity for the whole cardiovascular system is difficult to evaluate. The viscous loss of energy, measured at a cross-sectional level, although appearing small in absolute value, may be much higher if it is integrated all along the arterial tree.
In conclusion, active mechanisms, including normal endothelial function, appear to compensate for the intrinsic viscosity of the arterial wall in physiological conditions, thereby improving the efficiency of the cardiovascular coupling energy balance.
Selected Abbreviations and Acronyms
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Acknowledgments
This study was supported by a grant from the Société Française d’Hypertension Artérielle, which is part of the Société Française de Cardiologie, and from Institut National de la Santé et de la Recherche Médicale (INSERM grant No. 494014). We thank Brigitte Laloux for her excellent technical assistance.
Footnotes
Reprint requests to Stéphane Laurent, MD, PhD, Service de Pharmacologie, Hôpital Broussais, 96, Rue Didot, 75674 Paris Cedex 14, France.
Received March 5, 1996; accepted November 1, 1996.
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