The goal of this study is to
quantitatively describe the remodeling of the zero-stress state of the
femoral artery in flow overload. Increased blood flow, approximately as
a unit step change, was imposed on the femoral artery by making an
arteriovenous (a-v) fistula with the epigastric vein. The a-v
fistula was created in the right leg of 36 rats, which were divided
equally into six groups (2 days and 1, 2, 4, 8, and 12 wk after the
fistula). The vessels in the left leg were used as controls without
operative trauma. The in vivo blood pressure, flow, and femoral outer
diameter and the in vitro zero-stress state geometry were measured. The in vivo shear rate at the endothelial surface increased approximately as a step function by ~83%, after 2 days, compared with the control artery. The arterial luminal and wall area significantly increased postsurgically from 0.15 ± 0.02 and 0.22 ± 0.02 mm2 to 0.28 ± 0.04 and 0.31 ± 0.05 mm2, respectively, after 12 wk. The wall thickness did not
change significantly over time (P > 0.1).
The opening angle decreased to 82 ± 4.2 degrees postsurgically
when compared with controls (102 ± 4.4) after 12 wk and
correlated linearly with the thickness-to-radius ratio. Histological
analysis revealed vascular smooth muscle cell growth. The remodeling
data are expressed mathematically in terms of indicial functions, i.e.,
change of a particular feature of a blood vessel in response to a unit
step change of blood flow. The indicial function approach provides a
quantitative description of the remodeling process in the blood vessel wall.
residual stress; opening angle; flow-dependent dilation; arteriovenous fistula; indicial response functions
 |
INTRODUCTION |
RESIDUAL STRAIN IS
DEFINED as the strain in the no-load state (where external forces
are zero) in reference to the zero-stress state. It is well known that
residual strain exists in blood vessels and that the arterial
configuration at zero-stress state is an open sector (5,
31). The open sector can be characterized by an opening angle,
which is a measure of the residual strain in the arterial wall
(20). It has been shown that the residual strain reduces
the stress concentration at the inner portion of the arterial wall at
physiological loading (1). It was further pointed out that
a change of the opening angle is a result of nonuniform growth of
dimension of the tissues in the vessel wall (8).
There is no doubt that significant remodeling of the arterial blood
vessel occurs in response to flow overload. Although there are many
ways to study the structural and mechanical remodeling of the vessels,
we chose residual strain as the most relevant quantitative aspect of
remodeling because it is a measure of the nonuniformity of growth or
resorption in different parts of the blood vessel wall
(8). If growth or resorption were uniform in the vessel
wall, then the structural change will cause no change in stress and
strain in the blood vessel. On the other hand, if one part of the
vessel wall outgrows the rest (or resorbs locally), then it will
compress the remainder and cause internal stress. The internal stress
is hence a measure of the unequal growth or nonuniform remodeling. This
internal stress is called residual stress. Measurements of the residual
strain yield a simple index for the mismatch of hypertrophy and
hyperplasia in different parts of the vessel. Recent studies have
demonstrated that the residual strain changes reflect the nonuniformity
of growth and remodeling in hypertension, diabetes, aging, and smoke
exposure (16, 19, 20, 25). To our knowledge, the effect of
flow overload on the residual strain has not been studied. Therefore,
the first goal of the present study is to determine the remodeling of
the zero-stress state of the femoral artery in response to flow
overload. An approximate step increase in arterial blood flow was
created in rat femoral artery by an arteriovenous (a-v) fistula. The
arterial morphology and opening angle were measured, and the residual
strains at the inner and outer wall surfaces were computed at the
no-load state. Furthermore, the shear rate at the inner wall and
the average circumferential stress and midwall strain were computed in
the in vivo state.
A systematic approach to describe the remodeling process is to record
the course of changes that take place in various features of the blood
vessel after a step change of input variable. Response to a step
function is called the indicial response function (IRF), i.e., change
of a particular feature of a blood vessel (e.g., wall thickness,
opening angle, strain, stress, and so forth) in response to a unit step
change of input variable (e.g., flow). If the system is linear, then a
convolution of the IRF yields the response to any arbitrary input
function within the same limits. This approach was first proposed by
Fung (personal communication) and was later used by Liu and
Fung (18) to obtain the IRF of arterial remodeling in
response to changes in blood pressure. The IRFs of arterial remodeling
in response to changes in blood flow have not been determined. Hence,
the second goal of the present study is to obtain the IRF of various
features of blood vessel in response to flow overload.
| |
MATERIALS AND METHODS |
Indicial Response Experiments
Ideally, indicial responses should be measured in one vessel.
This is not possible, however, if histological measurements have to be
made at different stages of tissue remodeling in different animals.
Furthermore, there is no method to change blood flow as a strictly
mathematical step function. In practice, one has to use many animals
with approximate steps of flow and then analyze the results by methods
discussed below.
Experimental design.
Thirty-six female Wistar rats weighing 236 ± 36 g
(means ± SD) were used in this study. The rats first underwent an
end-to-side anastomosis of the epigastric vein to the femoral artery in
their right leg. The epigastric venous bypass increased blood flow in the femoral artery. The left leg was used as a control without surgical
trauma. The rats were randomly divided into six chronic groups that
survived 2 days and 1, 2, 4, 8, and 12 wk after the a-v fistula.
The experiments complied with the animal welfare regulations formulated
by the Danish National Society for Medical Research and the Guide
for the Care and Use of Laboratory Animals.
Surgical procedure.
Anesthesia was induced and maintained with Hyponorm (0.3 ml/kg im) and
Dormicum (0.4 ml/kg ip). Microsurgery was performed under sterile
conditions with the use of a Zeiss operating microscope. The femoral
artery and vein and the epigastric vein were dissected along their
emergence from the fat pad to the junction with the femoral vein. The
femoral artery was then dissected out of its sheath over its entire
length. The adventitia of the femoral artery was thoroughly removed
over an area twice as long and twice as wide as the arteriotomy. Two
cuts were made from opposite directions at 45 degrees to the vessel
such that they met exactly and formed a mouth "V." The epigastric
vein was cut at a sufficiently distal position to reach the femoral
artery. The epigastric venous lumen was washed by heparinized saline
solution (50 EU/ml). The end of the distal epigastric vein was sutured
to the femoral arterial side with the use of a 10/0 microvascular
monofilament nylon suture (S & T Marketing). The skin incision was
closed in one layer.
In vivo measurements.
At a scheduled time, the original incisions were reopened under a
standard anesthetic regime. The outer diameter of the femoral artery in
both legs was videotaped under a stereomicroscope (Zeiss, Stemi 2000-C)
and later measured with the use of an image analysis system (Optimas).
Short segments of the femoral vessels of both legs were isolated, and
the arterial blood flow rates were measured with the use of an
ultrasound flowmeter (T206 small animal blood flowmeter, Transonic
System). The systemic arterial blood pressure was recorded from the
left carotid artery (Cardio-Med System). We did not routinely measure
arterial pressure in the femoral artery to prevent damage to the short
femoral artery. In pilot experiments, however, it was verified that the
pressure in the femoral artery is similar to that in the carotid artery.
The rats were killed with an overdose of anesthesia. The vessels were
immediately excised and placed in an organ bath containing Krebs
solution (with EGTA, 100 mg/l) aerated with 95% O2-5%
CO2 at room temperature. Adjacent tissue was carefully
removed with the aid of a stereomicroscope.
No-load and zero-stress state of the femoral artery.
Three to four segments ~0.2 mm in length were cut from the proximal
femoral artery near the site of the anastomosis. The cross section of
the segmental rings in the no-load state (zero pressure) was
videotaped. The luminal area, wall area, inner and outer wall circumferential length, and wall thickness were measured with the use
of an image analysis system. The measurements were averaged over the
three to four rings and used for the subsequent analysis. The rings in
the no-load state were cut radially at the anterior position, which was
marked with ink. The radial cut causes the ring to open up into a
sector, which releases the residual stress and has been shown to
represent the zero-stress state (5). The cross sections of
sectors were videotaped 30 min after the radial cut to allow the vessel
creep to subside. The circumferential lengths of the inner and outer
surfaces were measured. The opening angle was defined in accordance
with Fung (5), i.e., the angle between two radii joining
the midpoint of the arc of the inner wall of the vessel to the tips of
the sector. All femoral arterial sectors were fixed in formalin for
histological preparation.
Histological preparations.
The vascular specimens were fixed in formalin over 24 h and
embedded in paraffin. Five-micrometer-thick sections were cut and
stained with hematoxylin and eosin and elastin-Trichrom stain. The smooth muscle, collagen, and elastin in the arterial wall were
distinguished with three different colors in the elastin-Trichrom stain. Cross-sectional wall area, vessel thickness, and smooth muscle
cell count per area were determined from the 8-bit BMP color images.
Biomechanical analysis.
Analysis of deformation requires calculations of strain. The
circumferential strain and stress of the arteries were determined under
the assumptions that 1) the material in the artery is
homogenous, and 2) the vessel shape is cylindrical. The
circumferential stretch ratio of the artery (
) can be
computed with respect to the zero-state state according to the
following equation
|
(1)
|
where l is the circumferential length
at the no-load or in vivo state, and L is the
circumferential length at the zero-stress state. In large deformation,
one should use Green's strain, defined as
|
(2)
|
where E is the circumferential strain
in either the no-load (residual strain) or loaded state (in vivo
strain) corresponding to the circumference
(l) at no-load or in vivo states,
respectively, in reference to the circumference
(L) at zero-stress state. Residual strains
were computed at the inner and outer wall surfaces, whereas in vivo
strain was calculated at the midwall. The midwall circumferential
length of the artery was calculated as the average between the inner
and outer circumferences.
At an equilibrium condition, the average circumferential stress in an
arterial wall at the in vivo pressure can be computed with an
assumption that the shape of the artery is cylindrical as
follows
|
(3)
|
where 
is the circumferential stress, P is the
inflation pressure, ri is the arterial inner
radius, and h is the wall thickness at the in vivo pressure.
With an assumption that material in the artery is incompressible, the
inner radius of the artery at the in vivo pressure can be computed on
the basis of the measured outer diameters at the in vivo and no-load
states and the stretch ratio in the longitudinal direction as given by
the incompressibility condition
|
(4)
|
where ro and ri
are the outer and inner radii at the in vivo state, respectively, and
Ao and z are the cross-sectional area at the
no-load state and axial stretch ratio, respectively.
The shear rate due to blood flow on the inner wall of the artery is
computed by the equation for laminar flow (
) in circular cylinders
|
(5)
|
where ri is internal radius of blood
vessels, and Q is the volumetric blood flow rate. Under the condition
of high shear rate, blood viscosity is approximately constant. The wall
shear stress is the product of the wall shear rate and the viscosity of blood.
Statistical analysis.
The results are expressed as means ± SE. The time course of the
changes in the artery after the creation of the a-v fistula and
differences between groups were examined with a two-way ANOVA. Student's t-test was also used to detect possible
differences between groups of data. The results were regarded as
significant if P < 0.05.
Analysis of Indicial Response Experiments
Because the experiments are done by approximate step change of
blood flow, we have to transform the results to those of exact steps by
computation. This can be done if the system is linear with respect to
the amplitude of the disturbance about a homeostatic state. Assuming
linearity, this can be done as follows. Let L(t) be a parameter of the system, e.g., the wall thickness, the
circumference, the opening angle, and so forth; t is
time. Let FLQ(t) be the IRF of
L(t) 
L(0) to a
unit step change of 
Q(t) = Q(0)H(t), where
H(t) is the heavyside step function,
the value of which is zero when t < 0, one when
t > 0, and one-half when t = 0. If the
flow perturbations consist of a jump from 0 to Q(0) at
t = 0 and a continuous function Q(t) that
has a finite derivative at t > 0, then under the
linearity assumption (see Fung, Ref. 7), the following
convolution integral applies
|
(6)
|
To determine the IRF, we used the method of Laplace
transformation. The Laplace transform of
Q(t) is 
(s), defined by the
integral
|
(7)
|
Similarly, the Laplace transforms of
L(t) and
FLQ(t) are obtained by multiplication
with e
st and integration of the
product from zero to infinity to obtain 
(s) and
(s). The Laplace
transformation of Eq. 6 is
|
(8)
|
Our experiments show a blood flow history that is equal to the
homeostatic flow plus a step and a perturbation
|
(9)
|
then
|
(10a)
|
and
|
(10b)
|
where A, A1,
A2, and A3 are empirical
constants determined from curve fitting of experimental data. The
response of the various parameters measured can be expressed in one of
the following three forms
|
(11a)
|
|
(11b)
|
|
(11c)
|
where B, b, B1,
b1, B2, and
b2 and C1,
C2, and C3 are empirical
constants determined from curve fitting of experimental data. The
Laplace transform of the input function, Q(t), is
|
(12)
|
The Laplace transformation of the output quantities are
|
(13a)
|
|
(13b)
|
|
(13c)
|
The Laplace transform of the IRF,
(s), can be
determined for each of the cases by substituting Eq. 12 and
the respective Eq. 13 into Eq. 8 as shown in the
APPENDIX. The inverse transformation of
(s) can then be
computed to obtain the IRF of remodeling for the various vessel
parameters (see APPENDIX).
| |
RESULTS |
In Vivo Data
The rats experienced small weight loss 2 days after the a-v
fistula and gained weight during the remaining study period
(P < 0.01), as shown in Fig.
1A. They gained ~30 g
(
12% of body wt) during the 3-mo period.
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Fig. 1.
In
vivo measurements. A: body wt. B: arterial blood
pressure in carotid artery. C: blood flow in the right and
left femoral arteries. D: outer diameter in the right and
left femoral arteries. R and L, right and left, respectively. The
anastomosis was performed on the right femoral artery, and the left
femoral artery was used as a control. a-v, Arteriovenous. Values are
means ± SE. C and D: * statistical
significance.
|
|
The systemic arterial blood pressure measured in the carotid artery is
shown in Fig. 1B. The arterial pressure did not change postoperatively (P > 0.05). The a-v fistula
created a significant increase in pressure drop along the femoral
artery because of the decrease in downstream pressure at the site of
the anastomosis. Hence, the blood flow rate in the anastomosed femoral
artery increased to a value three to four times larger than that of the
control leg 2 days after the a-v fistula (P < 0.01) and did not significantly change during the remaining
period (P > 0.5) (Fig. 1C). The
greatest increase in blood flow occurred at 2 wk after the fistula. The flow-induced remodeling of the in vivo outer diameter of the
femoral artery is shown in Fig. 1D.
No-Load and Zero-Stress State Data
The inner and outer wall circumferences at no-load and zero-stress
states are shown in Fig. 2, A
and B, respectively. Increased circumference indicates
arterial dilation. Both the luminal and wall area in the no-load state
increased postoperatively (P < 0.001) (Fig.
2, C and D, respectively). This was further
evidenced by the histological findings as will be discussed (see
Histological Data). The wall thickness did not differ
between the control and fistula groups (P > 0.1). There was also no statistically significant increase
in wall thickness with time in control arteries (P > 0.1) (Fig. 2E). The thickness-to-radius ratio,
however, decreased significantly with time compared with the control
group (P < 0.05), as shown in Fig.
2F. This was due to an increase in lumen radius. All data of
control arteries (Fig. 2, A-F) showed no
statistically significant variation with time (P > 0.5). With the exception of wall thickness
(P > 0.5), all other morphometric
parameters (Fig. 2, A-F) showed significant
changes after the creation of the a-v fistula (P < 0.001).
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Fig. 2.
Ex
vivo measurements. A: inner circumference at the no-load
(nl) and zero-stress (zs) states in the right and left femoral
arteries. The differences in the inner circumference at the no-load and
zero-stress states between the right and left femoral arteries were
statistically significant at all time intervals  2 wk and 4 wk,
respectively. B: outer circumference at the no-load and
zero-stress states in the right and left femoral arteries. The
differences in the outer circumference at the no-load and zero-stress
states between the right and left femoral arteries were statistically
significant at all time intervals 2 wk and 8 wk, respectively.
C: lumen area at the no-load state in the right and left
femoral arteries. D: wall area at the no-load state in the
right and left femoral arteries. E: wall thickness at the
no-load state in the right and left femoral arteries. F:
wall thickness-to-radius ratio at the no-load state in the right and
left femoral arteries. The anastomosis was performed on the right
femoral artery, and the left femoral artery was used as a control.
Values are means ± SE. C-F:
* statistical significance.
|
|
The temporal change of the opening angle is shown in Fig.
3A. The opening angle did not
show significant change in the control group (P > 0.1). In the a-v fistula group, the opening angle was unchanged after 4 wk and then subsequently decreased (P < 0.01) compared with the control group. The difference
between fistula groups (at 8 and 12 wk) and control groups was
statistically significant (P < 0.05). The
arterial residual strain at both the inner and outer surface is
presented in Fig. 3B. It can be seen that the outer strain
is tensile, whereas the inner strain is compressive for both the
control and fistula groups. The average outer residual strain becomes
more tensile in the 1- and 2-wk groups, whereas the average inner
residual strain becomes less compressive in the 4-, 8-, and 12-wk
fistula groups compared with the control. The temporal variations of
the inner and outer residual strains in the control and fistula groups
were not found to be statistically significant (P > 0.1). Figure 3C shows a positive correlation between the opening angle and the wall thickness-to-radius ratio (r = 0.906, P < 0.05).
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Fig. 3.
A: opening angle of the right and left femoral
arteries. B: computed residual strains at the inner and
outer wall of the right and left femoral arteries. C:
correlation of opening angle and wall thickness-to-radius ratio of
femoral artery. The correlation coefficient for the linear fit is 0.90 (P < 0.02). Values are means ± SE.
A and B: * statistical significance.
|
|
Biomechanical Data
The incompressibility condition expressed in Eq. 4 was
used to compute the inner radius of the femoral artery, as shown in Fig. 4A. It can be seen that
the inner radius in the femoral artery increased from an average value
of 0.33 to 0.40 mm in 12 wk after the creation of the fistula. We also
computed various biomechanical parameters including wall shear rate,
midwall circumferential strain, and mean stress, as shown in Fig. 4,
B-D, respectively. After the creation of the
a-v fistula, shear rate, midwall strain, and mean stress significantly
increased initially and decreased thereafter (P < 0.001).
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Fig. 4.
Computed parameters at the in vivo state. A: inner
radius of right and left femoral arteries. B: shear rate at
the inner wall surface of right and left femoral arteries.
C: midwall strain in the right and left femoral arteries.
D: mean circumferential stress in the right and left femoral
arteries. The anastomosis was performed on the right femoral artery,
and the left femoral artery was used as a control. * Statistical
significance.
|
|
Histological Data
In control arteries, the tunica media is composed of an
average of four to five smooth muscle cell layers. It is bounded by a
heavily stained, single-layered internal and external elastic lamina.
There are several layers of continuous elastic fibers throughout the
tunica media. The tunica intima consists of a confluent endothelial
monolayer. Figure 5 shows a
photomicrograph of a right femoral artery during the progression of
flow overload. The relationship between the cross-sectional area of
arterial media in the zero-stress state and the duration of the a-v
fistula for the two groups is shown in Fig.
6A. The change in the medial
area of the control artery throughout the 12-wk period was not
statistically significant. At 4 wk after fistula, the femoral medial
area was increased compared with control. The increase in medial
cross-sectional area was due to both net smooth muscle cell (SMC)
proliferation and cell hypertrophy. Figure 6B shows the SMC
proliferation during the progression of the fistula. The change in
the SMC count (the total number of SMC in the medial wall of the
open sector, i.e., in the zero-stress state) of the control artery was
not statistically significant in the 12-wk period. After 4 wk of flow
overload, the SMC count was significantly larger in the fistula
group compared with the control group. The large increase in
medial cross-sectional area, however, was mainly due to SMC
hypertrophy. Small increases in intimal thickness were observed in
some parts of some arteries.
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Fig. 5.
Photomicrographs of right femoral artery cross sections
in control artery and 2 and 12 wk after anastomosis (left to
right, ×400 magnification).
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Fig. 6.
Histological measurements at the zero-stress state. A:
area of media in the open sector of the right and left femoral
arteries. B: the total number of smooth muscle cells (SMC)
in the media of the open sector of right and left femoral arteries. The
anastomosis was performed on the right femoral artery, and the left
femoral artery was used as a control. Values are means ± SE.
* Statistical significance.
|
|
IRF Data
The constants A, A1,
A2, and A3 in Eq. 9, characterizing the change in blood flow, were found to have
values of 18.6 ml/min, 9.38 ml/min, 0.214 ml · min1 · wk1, and
0.00288 ml · min1 · wk2,
respectively, as determined by a least square fit of the flow data
(R2 = 0.982). The exponential function,
Eq. 11a, was used to fit the experimental data for inner and
outer circumferences, inner wall area and thickness-to-radius ratio in
the no-load state, inner circumference at the zero-stress state, and
internal diameter in the in vivo state. The least square fit constants
B and b along with the correlation coefficient
are listed in Table 1 for the various
parameters. Similarly, a biphasic function, Eq. 11b, was used to fit the experimental data on the inner and outer strain in the
no-load state and flow shear rate and the midwall strain in the in vivo
state. The empirical constants are given in Table 1. Finally, a cubic
function, Eq. 11c, was used to fit the data on the outer
circumference and opening angle in the zero-stress state and the wall
area in the no-load state with the corresponding constants listed in
Table 1.
The curves showing the time history of various parameters (e.g., Figs.
2-4 and 6) are not indicial curves, because the flow change,
although having an approximate step increase at t = 0, did not remain constant at t > 0. To simplify the
interpretation, we need to extract the IRFs from these curves. Figures
7-9 are the IRFs computed from the
experimental data shown in Table 1 with the formulas given in the
APPENDIX (Eqs. 20, 24, and 27 for
cases I, II, and III, respectively).
The IRFs of inner and outer vessel circumferences in the no-load and
zero-stress states are shown in Fig. 7, A and B,
respectively. Figure 7C shows the indicial response of the
opening angle. The histological data on the area of media and SMC count
were also extracted in the form of IRFs as shown in Fig.
8, A and B,
respectively. Finally, the various biomechanical parameters (radius,
shear rate, circumferential strain, and stress) were also expressed in
terms of IRFs as shown in Fig. 9,
A-D.
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Fig. 7.
Indicial response functions (IRF). A: inner
and outer circumference in the no-load state. B: inner and
outer circumference in the zero-stress state. C: opening
angle.
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Fig. 8.
IRFs. A: medial area in the open sector of the femoral
artery. B: total number of SMC in the media of the opening
sector of the femoral artery.
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Fig. 9.
IRFs. A: inner radius. B: wall shear rate.
C: inner and outer circumferential (Circ) residual strain
and in vivo midwall strain. D: mean circumferential
stress.
|
|
| |
DISCUSSION |
Flow-Induced Remodeling of Blood Vessels
The first author to clarify the relationship between blood flow
and blood vessel remodeling was Thoma (30), who in 1892 observed that in chick embryos, certain pathways of most rapid blood
flow increased in caliber and length. Thoma observed that the growth of
blood vessel lumen depends on the flow, and the wall thickness depends
on the tension in the wall. Schretzenmayr (26) confirmed
Thoma's observations in 1933. More recent studies have shown that
increased blood flow induces blood vessel dilatation even in small
muscular arteries (15, 29). Flow-induced dilation is found
to be influenced by local endothelial cell function (15, 28). It is now accepted that shear stress acts through the
endothelium to regulate both acute vessel tone and chronic remodeling
of blood vessel (2). The endothelium acts as a complex
mechanical signal-transduction interface between the flowing blood and
the vessel wall (2, 23).
Uniform Shear Hypothesis
The constant wall shear rate hypothesis implies that the
volumetric flow rate is proportional to the cube of the vessel radius, assuming a laminar, steady-state, and incompressible Newtonian flow
through a rigid cylindrical vessel (13). Hence, to
maintain a constant wall shear rate, the cube of the radius must
increase in proportion to the increase in blood flow. Blood vessels can accommodate such a change in vessel radius at two levels: acutely through vasoactive mechanisms (flow-dependent constriction or dilation)
and chronically by adjusting vascular caliber (12, 32).
A number of studies have previously shown a normalization of wall shear
strain after a significant increase in blood flow (10-12,
32). Kamiya and Togawa (12) created an a-v fistula between the common carotid artery and the external jugular vein in dogs
and studied the remodeling 6-8 mo postoperatively. They found a
normalization of shear rate to within 15% despite a fourfold increase
in blood flow. Additional a-v fistula studies of the iliac artery of
the monkey and rabbit (21, 32) and radial artery of humans
with end-stage renal disease (10) also showed a
normalization of shear rate by remodeling of vascular caliber. In the
present study, we increased the blood flow in the rat femoral artery by
creating an epigastric venous-to-femoral arterial fistula, which
subsequently increased the wall shear rate or stress (if viscosity is
constant) by ~83% after 2 days. The shear rate decreased to within
27% of the control value, however, after a 12-wk period. The
differences in shear rate at 12 wk were not statistically significant,
as shown in Fig. 4B.
Remodeling of the Zero-Stress State
Mechanical factors have been proposed to regulate growth and
remodeling of biological tissue (6). Since Fung
(5) and Vaishnav and Vossoughi (31) found
that residual stress exists in the arterial wall when external loads
are absent (no-load state), the zero-stress state has been used as a
reference state for mechanical analysis. Furthermore, it is best to
measure the structural components of the vessel wall at the zero-stress
state, because in this state the morphology and sizes of the cells and
extracellular matrix are not distorted by stress and strain. Otherwise,
there is the complication of deformation due to internal stress
(6). Residual stress can be altered by many factors,
including tissue growth, remodeling (changes in material properties),
and acute geometrical changes (22). Experimental evidence
has shown that the opening angle of an artery increases during
hypertension, which indicates an increase in the residual strain
(9, 20). The decrease in the opening angle in response to
flow overload shown in this study suggests that the loading pattern is
an important determinant of growth and remodeling of the zero-stress state.
Experimental and theoretical evidence suggests that geometric
remodeling alters the residual strain in the vessel wall
(22). In a theoretical analysis, Rodriguez et al.
(24) predicted that concentric hypertrophy, which
increases the wall thickness-to-radius ratio, increases the opening
angle and hence the residual stress. They also showed that eccentric
hypertrophy characterized by a decrease in the wall thickness-to-radius
ratio may induce a reversed transmural residual stress gradient. Omens
et al. (22) provided experimental evidence for this
prediction in ventricular remodeling during cardiac postnatal growth in
rats. Their study showed that eccentric hypertrophy decreased the
opening angle and residual stress in the circumferential direction.
Those results in the heart are consistent with our flow-overload
remodeling in the femoral artery. This is in contrast to concentric
remodeling induced by hypertension, where the opening angle and the
residual stress are increased. In hypertension, it is proposed that
the increase in residual stress makes the transmural stress and
strain distributions more uniform (20). Omens et al.
(22) suggested that the decrease in opening angle with a
decrease in wall thickness-to-radius ratio in response to flow overload
occurs for purely geometric reasons, because the vessel becomes a
thin-walled cylinder having a more uniform transmural stress
distribution independent of residual stress.
In the present study, the decrease in opening angle is consistent with
Fung's (8) hypothesis of nonuniform remodeling; i.e., if
the inner wall grows more than the outer wall, the opening angle will
be increased, whereas if the outer wall grows more than the inner wall,
the opening angle will be decreased. Figure 7 shows that in flow
overload, the outer wall grows more than the inner wall, and hence the
opening angle is decreased.
Remodeling of Circumferential Stress and Strain
An increase in wall shear stress causes circumferential dilation,
which increases the circumferential stress and strain in the vessel
wall. This implies that circumferential stress and strain are
additional factors that may play a role in arterial wall remodeling in
flow overload. Observation on vein grafts provides evidence to support
this hypothesis. Dobrin et al. (3) constrained the vein
grafts circumferentially by partially enclosing the grafts in a cuff.
The vein grafts were exposed to arterial pressure and flow. The part of
the graft located in the cuff could not dilate, whereas the other part
could dilate. The two parts were exposed to the same arterial pressure,
but the circumferential stress and strain were different. Their results
showed that the increases in the vessel medial thickness and diameter
were best associated with increased wall stress and strain in the
circumferential direction. Other observations from vein grafts suggest
the same conclusion (14, 27). Hence, circumferential
deformation may play a role as a regulatory factor in arterial
remodeling caused by flow-induced dilation.
The previous studies on vascular remodeling in response to physical
stress (3, 10-12, 21, 32) did not assess the
circumferential strain because the reference state, zero-stress state,
for the evaluation of strain was not investigated. One of the goals of the present study was to characterize the remodeling of the zero-stress state and to use the results to compute the strain. In our a-v fistula
model, the midwall strain and circumferential stress were increased by
~53 and 17%, respectively, after 2 days. The vascular remodeling
occurred in such a way as to normalize the midwall strain and
circumferential stress of the vessel wall, which were found to be
within 22 and 12% of their respective control values at the 12-wk
period. These differences were not statistically significant, as shown
in Fig. 4, C and D. In fact, the differences between the fistula and control groups were not statistically significant beyond 4 wk and 8 wk for the mean stress and midwall strain, respectively.
The physical factors (wall shear rate, circumferential stress, and
midwall strain) that stimulate vascular remodeling may be expressed in
terms of the IRFs for the three biomechanical parameters as shown in
Fig. 9, B-D. To compare the relative
magnitude of the three biomechanical IRFs, we normalized the quantities with respect to their initial values as shown in Fig.
10. It can be seen that the two major
biomechanical stimuli for remodeling are wall shear strain and midwall
strain. The midwall circumferential strain has a shorter rise time and
a shorter decay time than the wall shear rate. This suggests that the
circumferential strain may be an important initial stimulus for
remodeling.
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Fig. 10.
IRFs of shear rate, mean circumferential stress, and
midwall (Mid) strain normalized to the respective initial values.
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IRFs
A systematic approach to describe the remodeling process is to
extract the IRFs as given by Eqs. 20, 24, and
27. The use of IRFs simplifies the interpretation of data
and greatly increases the potential of using experimental data for
prediction of the outcome of future experiments under an arbitrary
course of stimulation. It also allows the reduction of complicated sets
of experimental data into definitive statements in terms of IRFs.
Finally, it presents a quantitative way to verify the basic hypothesis
of linearity. This is a powerful engineering approach that provides a
rigorous method to study the function of blood vessels by properly formulated boundary-value problems and makes possible a wide variety of
applications to the study of tissue engineering.
Critique of Methods
The surgical trauma (including mechanical dissection, compression,
ischemia, barometric and osmotic changes, and so forth) involved in the construction of an a-v fistula may stimulate vascular remodeling and hyperplasia. Because a sham operation was not done on
the control leg, this raises the issue of whether any of the biomechanical changes are a result of surgical trauma. Although we did
not fully address this issue, we did obtain data that suggest that the
effect of surgical dissection is small. In five animals, both the left
and right femoral arteries were dissected in preparation for the a-v
fistula. The fistula, however, was only created in the right femoral
artery. Hemodynamic and morphological data were obtained at 8 wk (2 animals) and 12 wk (3 animals) after the initial surgery. We found no
statistically significant differences between the dissected left
femoral artery and the corresponding nondissected control in the 8- and
12-wk groups.
The equation used to compute the shear rate, Eq. 5, is based
on the assumption that blood flow has a Poiseuille profile. To examine
this hypothesis, at the site where the morphometric measurements were
made, we must examine the architecture of the vessel. The femoral
artery arises from the external iliac artery. Approximately 3 mm from
the external iliac bifurcation, the femoral artery gives rise to a
Morphy branch, which is 0.07 mm in diameter. The measurements were made
at 1 mm downstream of the Morphy branch. The anastomosis was ~3 mm
downstream of the site of measurements. The issue here is as follows:
Does the entry flow into the femoral artery become fully developed at
the site of measurements? In other words, what is the inlet length
(which is defined as the distance through which the velocity profile
becomes approximately parabolic)? Lew and Fung (16) have
previously shown that the entrance length (Le)
is approximately given by Le = 0.08ReD, where
Re and D are the Reynolds number and
diameter of the vessel, respectively. This expression applies for
Reynolds numbers in the range of 50-100. The Reynolds number is
given by Re = (4
Q)/(
µD),
where , µ, and Q are the density, viscosity, and flow rate of
blood, respectively. The range of Reynolds numbers was found to be
48-120 for the range of flow rates and diameters (5.0-18.2
ml/min and 0.60-0.80 mm, respectively). Hence, we obtain an
entrance length of 2.3-7.6 mm. Because the diameter of the Morphy
branch is relatively small, there is an effective length of 4 mm from
the external iliac artery to the site of measurements. Hence, the
velocity profile should be nearly parabolic, and Eq. 5 is a
reasonable approximation.
Future Studies
Equations 20, 24, and 27
are based on the linearity assumption of the response function
L(t) with respect to the amplitude of the flow
step. A system is linear if L(t) is proportional
to Q(0) in Eq. 6. Otherwise, the system is
nonlinear. The linearity of the system can be tested in two ways.
First, one can directly test the relationship between
L(t) and Q(0) by varying the
magnitude of the step function Q(0) and measuring the
corresponding values of L(t). Experimentally, the
flow rate through the fistula can be varied by constriction of the
epigastric vein. Second, validation of linearity can be done by
performing a new perturbation, predicting the results according to
Eq. 6, and comparing the predication with measured results.
For example, one can impose a flow history by constriction of the
epigastric vein with an ameroid. The ameroid will slowly occlude the
epigastric vein and hence yield a monotonic decrease in flow after the
initial rise. The predictions of Eq. 6 can be compared with
the experimental results of the ameroid experiment to determine the
limits of validity of the linearity hypothesis.
If a system is nonlinear, but a small change in Q(0)
results in a small change in L(t), then Eq. 6 can still be used. If a small change in Q(0)
produces a large change in L(t), then the system
is grossly nonlinear and Eq. 6 cannot be used. In this case,
a record of the way the indicial function depends on
Q(0) is the best way to quantitatively express the
nonlinearity; i.e., FLQ is not only a
function of t but is also a function of the magnitude of Q. Modifying the IRFs in this way, one can still use Eq. 6,
which now becomes nonlinear. The range of agreement between theory and
experiment will correspond to the range of linearity. The goal of
future experiments is to find the upper and lower limit to linearity.
The limits are related to the borderline between physiology and pathology.
TOP
ABSTRACT
INTRODUCTION
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, and 
are given by experimental results (Table 1),
FLQ(t) can be derived mathematically
from the experimental data by using Eq. 20. More complex
formulas needed to handle the experimental data are considered in
cases II and III.
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