Download Right coronary artery becomes stiffer with increase in elastin and

J Appl Physiol 106: 1338–1346, 2009.
First published January 29, 2009; doi:10.1152/japplphysiol.90592.2008.
Right coronary artery becomes stiffer with increase in elastin and collagen
in right ventricular hypertrophy
Marisa Garcia2 and Ghassan S. Kassab1
1
Department of Biomedical Engineering, Surgery, Cellular and Integrative Physiology, Indiana University Purdue University,
Indianapolis, Indiana; and 2Department of Biomedical Engineering, University of California, Irvine, California
Submitted 29 September 2008; accepted in final form 26 January 2009
stress; strain; wall shear stress; constitutive equation
BLOOD VESSELS ARE subjected to the mechanical loading of blood
pressure and blood flow throughout the cardiac cycle. The
vessels undergo structural and mechanical adaptations in response to changes in the local hemodynamic conditions. For
example, changes in blood flow cause changes in fluid wall
shear stress (WSS; the tractive frictional force exerted by
flowing blood on the inner layer of the vessel wall) by eliciting
a wide range of biochemical and physiological responses in
experimental (46) and clinical studies (16, 22). Studies have
demonstrated that a sudden increase in blood flow initiates an
enlargement of the arterial lumen that progresses for months as
the WSS tends to a homeostatic or normal level (23, 25). A
gradual increase in blood flow has been reported by several
groups to result in outward hypertrophic remodeling of resistance (4, 41, 44, 45, 47) and conductive arteries (25, 36, 39).
The flow-induced remodeling may be an attempt to maintain
mechanical homeostasis (24, 25). When blood flow is elevated
in a vessel, the endothelium responds to restore WSS toward
Address for reprint requests and other correspondence: G. S. Kassab,
Dept. of Biomedical Engineering, SL-174, Indiana Univ. Purdue Univ.
Indianapolis, 723 West Michigan St., Indianapolis, IN 46202 (e-mail:
[email protected]).
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homeostatic levels by changing vascular tone, increasing the
lumen diameter (44), and potentially altering the function and
mechanical properties of the vessel wall. A detailed analysis of
these adaptations is necessary to understand vascular function
and dysfunction, including atherogenesis (36, 37).
The increase in stiffness of arteries represents an early risk
factor for cardiovascular diseases (1, 19). Specifically, increased stiffness is associated with aging (14), atherosclerosis
(6), and heart failure (27). It is also associated with various risk
factors including hypertension (7), diabetes (42), hyperlipidemia (33), and smoking (43). Furthermore, arterial stiffness has
also been shown to be an independent risk factor for cardiovascular events such as primary coronary events (3). Therefore,
the assessment of coronary mechanical properties is important
in understanding the processes of coronary artery disease.
Recently, our group (37) characterized the mechanical properties of normal coronary blood vessels in the range of physiological loading. We used a method of analysis to compute the
biaxial (circumferential and axial) incremental elastic moduli
under in vivo (homeostatic) conditions to characterize the
stiffness of the vessel wall (37). Briefly, the method utilizes a
well-accepted approach to nonlinear elasticity that uses a
linearized incremental formulation. It was demonstrated that
the normal incremental moduli under homeostatic conditions
are layer (intima-media and adventitia) and direction (circumferential and axial) dependent.
The objective of the present study was to quantitatively
describe the remodeling of the right coronary artery (RCA) in
response to flow-overload during right ventricular hypertrophy
(RVH) in response to pulmonary artery banding. Our hypothesis is that the remodeled proximal RCA has increased homeostatic moduli in association with increases in elastin and
collagen in the vessel wall. Accordingly, we determined the
incremental elastic moduli in the circumferential, axial, and
cross directions in the physiological range and the corresponding changes in elastin and collagen of the vessel wall using
multiphoton microscopy (MPM) for the control and remodeled
vessels in RVH. It was found that the moduli are significantly
increased in all directions in line with an increase in elastin and
collagen of media but more prominently of the adventitia. The
physiological and pathophysiological implications of the increased stiffness in RCA are highlighted.
METHODS
Animal preparation. Ten male Yorkshire farm pigs were used in
this study (5 RVH and 5 control). RVH was induced through acute
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Garcia M, Kassab GS. Right coronary artery becomes stiffer with
increase in elastin and collagen in right ventricular hypertrophy.
J Appl Physiol 106: 1338 –1346, 2009. First published January 29,
2009; doi:10.1152/japplphysiol.90592.2008.—Changes in blood flow
influence the structure, function, mechanical properties, and remodeling of arteries. The objective of the present study was to investigate
the role of increased blood flow on the biaxial incremental elastic
moduli of the porcine right coronary artery (RCA) and to determine
the microstructural basis for the changes in moduli. We hypothesized
that an increase in RCA flow will lead to increased stiffness in
conjunction with remodeling of elastin and collagen in the vessel wall.
The control and experimental groups consisted of five RCA vessels
each. The RCA of the experimental group was exposed to 4 wk of
flow-overload in right ventricular hypertrophy induced by pulmonary
artery banding. Stress-strain relationships were determined and the
incremental elastic moduli were derived in the circumferential, axial,
and cross directions. The results show a significant increase in the
elastic moduli in the circumferential (262.7 ⫾ 15.7 vs. 120.2 ⫾ 12.4
kPa; P ⬍ 0.001), axial (177.8 ⫾ 25.5 vs. 100.3 ⫾ 11.9 kPa; P ⫽
0.025), and cross directions (104.8 ⫾ 8.2 vs. 68.2 ⫾ 8.6 kPa; P ⫽
0.016) of the experimental RCA compared with controls. Multiphoton
microscopy was used to assess the changes in elastin and collagen
content in the media and adventitia of the vessel wall. We found a
significant increase in elastin and collagen area fraction particularly in
the adventitial layer. These data suggest stiffening of the vessel wall
as a result of increased elastin and more predominantly collagen.
STIFFER RCA IN FLOW-OVERLOAD
J Appl Physiol • VOL
from 0 to 125 mmHg at every stretch ratio. The Biopac system was
used for continuous data acquisition. The vessels were preconditioned
to stretch and pressure several (⬃5–7) times to ensure repeatable
loading and unloading responses for both groups identically. After
preconditioning, the forces generated and deformations were continuously
acquired during loading. The total experimental duration was ⬍4 h.
Morphometric measurements. After being tested, the coronary
segment was removed from the luers and two rings, sufficiently
distant (1–2 mm) from the ligated ends, were cut (one from each end)
for the morphometric measurements of no-load and zero-stress states.
Both rings were allowed to stabilize for 1–2 h in saline solution at
room temperature (22°C). The rings were photographed in the no-load
state before a cut was made radially at the anterior position allowing
each sector to approach the zero-stress state. The rings opened into a
sector and gradually approached a constant opening angle defined as
the angle subtended by two radii connecting the midpoint of the inner
wall to the ends of the open segment. The cross section of each sector
was photographed 1 h after the radial cut. Using a morphometric
analysis system (Sigma Scan Pro.5.0), the inner and outer circumferences and area of the rings in the no-load state were determined. The
circumference of the border between the medial and adventitial layers
was determined for each no-load and zero-stress state ring. The
morphological measurements of inner and outer circumferences and
opening angle in the zero-stress state were also acquired.
Microscopic visualization of the vessel. The arterial wall contains
collagen and elastin fibers and smooth muscle cells as the primary
load-bearing components. MPM was used to visualize the microstructural components of excised porcine blood vessels through twophoton excited fluorescence (TPEF) and second-harmonic generation
(SHG) microscopy. The current protocol for the visualization of
elastin and collagen, exclusively, in porcine coronary arteries has been
validated by Zoumi et al. (49) and is briefly described here. A 1- to
1.5-mm-long ring was obtained from the remaining, untested RCA
and, in the zero-stress state, fixed in 10% formalin. An upright
microscope was utilized making it necessary to use a special thinglass-bottom chamber in which a hole was drilled in the center of a
culture dish and coverslipped to create a short well. The well was
surrounded by a thin ring of grease and the specimen was placed
inside with a small amount of saline (0.9%) solution at room temperature (22°C). A coverslip was placed on the grease and gently pressed
down until just touching the segment. The dish was overturned and
mounted on the upright microscope stage. A Zeiss LSM 510 Meta
confocal/MPM system equipped with a tunable Titanium-Sapphire
laser was used to scan the vessel thickness of arterial specimens. The
samples were imaged with a 40⫻ corr, C-Apochromat, numerical
aperture 1.2, water immersion microscope objective. Two-dimensional (x-y plane) images (512 ⫻ 512 pixels) were acquired from
various depths (z) into the vessel wall at a rate of 9 frames/s (pixel
dwell time of 1.6 ␮s/pixel). Resolution was ⬃0.45 ␮m and 0.5 ␮m in
the x-y and z image planes, respectively. MPM images were obtained
using a 385– 425 emission filter for SHG and a 480 –520 band-pass
infrared (IR) emission filter for TPEF. Each acquired image was
typically integrated over a few seconds. The scan time for an entire
vessel wall was ⬍1 h.
Histology. Tissues were fixed (10% formalin) for at least 48 h
before processing. Five rings from each group, control and experimental, were rinsed with a buffer solution, dehydrated, and embedded
in paraffin. The 5-␮m-thick sections were mounted on glass slides and
stained with Verhoeff’s Elastic stain. The nuclei counts were performed in the media using the ImageJ software.
Statistical analysis. Student’s t-test was used to detect differences
between control and experimental vessels. A P ⬍ 0.05 is indicative of
a significant difference between the population means. Values are
expressed as means ⫾ SE.
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pulmonary artery banding similar to Kassab et al. (26). All experiments were performed in accordance with national and local ethical
guidelines, including the Institute of Laboratory Animal Research
(ILAR) Guide, Public Health Service policy, Animal Welfare Act, and
institutional policies regarding the use of animals in research and an
approved Indiana University Purdue University IACUC protocol.
Ketamine hydrochloride (20 mg/kg im) and atropine (0.05 mg/kg
im) were used to preanesthetize the animals. Endotracheal intubation
was performed and surgical anesthesia was mechanically maintained,
via a face mask, through the inhalation of isofluorane (0.5–2%) and
oxygen. An incision was made on the neck to chronically implant 7Fr
sheaths in the left carotid artery, for arterial pressure measurement,
and in the right external jugular vein, for drug administration. Lidocaine (80 mg iv) was administered as a bolus before cardiac instrumentation. A left lateral fourth-intercostal-space thoracotomy was
performed using sterile techniques. A small segment of the proximal
RCA was dissected free of surrounding tissue and a flow probe
(Transonic) was placed around it to measure coronary flow. A glycerin-filled occluder was placed around the pulmonary artery and the
associated tubing was exteriorized for inflation of the occluder after
recovery. The thoracotomy was closed, and the animal was allowed to
recover for 1 wk.
After 1 wk of recovery, the animal was preanesthetized and
surgical anesthesia was maintained, as described above, while a
fluid-filled silastic pressure-monitoring catheter was inserted through
the jugular sheath and advanced into the right atrium and right
ventricle. The degree of banding of the pulmonary artery (PA) was
gauged from the systolic RV pressure. The occluder was slowly filled
with glycerin which stenosed the PA until the RV systolic pressure
was raised by 9.3–10.7 kPa (70 – 80 mmHg). The occluder was then
fixed to maintain this degree of stenosis over a 4-wk period, during
which time the RV pressure was hypertensive, the cardiac muscle
hypertrophied, and the RCA flow increased.
After the 4-wk banding period, the animal was again preanesthetized and surgical anesthesia was administered as described above. A
midline sternotomy was performed and a bolus of normal saline with
200 U/kg of heparin was injected through the carotid artery catheter.
A set of angiograms was obtained for a separate study. The pigs were
then euthanized by an injection of pentobarbital sodium (300 mg/kg)
through the jugular vein. After harvest, the heart was placed in an
ice-cold saline bath. The degree of RVH was assessed by measuring
right-to-left ventricle weight ratio (8).
In situ preparation of coronary vessel. About 3.2 cm of proximal
RCA length was dissected free of fat and connective tissue under a
dissection microscope ⬃1–2 cm from its emergence at aortic ostia and
every branch was ligated to eliminate leakage from the blood vessel.
The vessel was marked in situ and the axial retraction upon excision
was documented to compute the axial stretch ratio (ratio of in
situ-to-unloaded length). By tying silk sutures (3-0) around the insertion sites, a luer was secured in place on either end of the vessel.
Ultrasonic crystals (Sonometrics) were sutured (10-0) diametrically at
the midpoint of the vessel segment. The changes in diameter were
continuously recorded with the crystals during mechanical testing.
Experimental measurement. The protocol for biaxial (inflation and
axial extension) testing of coronary arteries was recently described by
our group (37, 38). Here, we modified the set-up by the addition of
ultrasound crystals. Briefly, the arterial specimen with attached crystals was horizontally mounted on a TRIAX machine where the vessel
was kept in a saline bath during testing. The test set-up included a
pressure transducer and a manually driven linear stage where the axial
force was recorded by a load cell. The pressure transducer was
calibrated against a hydrostatic mercury column at pressures of 0 and
100 mmHg. The load cell was calibrated with a series of weights, and
a linear relationship was confirmed. During testing, the axial stretch
ratio (␭z; defined as the change in axial length between loaded and
no-load state) was varied over a small range (1.3–1.5) relative to the
in vivo loading. A pressure regulator was utilized to vary the pressure
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STIFFER RCA IN FLOW-OVERLOAD
RESULTS
Fig. 1. Representative example of outer diameter measurements with ultrasonic crystals during a pressure ramp in control and flow-overload right
coronary artery (RCA).
puted according to Eqs. 1b and 4, respectively (APPENDIX). The
circumferential and axial stresses and strains in conjunction
with Eqs. 5 and 6 (APPENDIX) yield the values of the incremental
moduli. The data on the circumferential, axial, and cross
incremental moduli for the control and experimental RCA are
summarized in Fig. 3. The results show a significant increase in
the moduli of elasticity in the circumferential (P ⬍ 0.001),
axial (P ⫽ 0.025), and cross (P ⫽ 0.016) directions of the
experimental vessels compared with the control group. Figure 4
compares the mean circumferential Cauchy stress and Green
strain for both groups at physiologic loading (pressure of 100
mmHg and stretch ratio of ␭z ⫽ 1.4). The Cauchy stress is seen
to be significantly elevated in the experimental group (P ⫽
0.032) while the Green strain remains unchanged (P ⫽ 0.949).
This finding reflects the restoration of circumferential strain at
elevated stress.
MPM was used to study the remodeling of the threedimensional structure of elastin and collagen of intact, unstained tissue. The specimen generated TPEF signals from
elastin and SHG signals exclusively from collagen for the
excitation wavelength used (800 nm) (49). Figure 5 illustrates
two-dimensional (230 ⫻ 230 ␮m) representative visualizations
of coronary arterial walls at 40⫻ magnification. The images
Table 1. Characteristics of RVH model
n
Body wt, kg
Heart wt, g
RV/(LV⫹S)
RV pressure, mmHg
Systemic BP, mmHg
RCA flow rate, ml/min
RCA diameter, mm
WSS, dyn/cm2
WT, mm
WTA, mm
WTI-M, mm
␭z
Control
RVH
5
43.4⫾5.0
220⫾22.0
0.31⫾0.01
31.2⫾0.88
64⫾6
39.9⫾4.03
3.61⫾0.17
10.7⫾0.55
0.48⫾0.043
0.20⫾0.017
0.28⫾0.025
1.40⫾0.00
5
43.7⫾1.90
252⫾24.7
0.73⫾0.093*
80.9⫾11.8*
68⫾8
81.3⫾13.1*
4.16⫾0.16*
12.5⫾1.84
0.46⫾0.028
0.20⫾0.012
0.26⫾0.015
1.43⫾0.01*
Values are means ⫾ SE; n ⫽ no. of pigs. RVH, right ventricular hypertrophy; RV, right ventricle; LV, left ventricle; S, septum; BP, blood pressure;
WSS, wall shear stress; WT, vessel wall thickness for right coronary artery,
RCA; WTA, adventitial layer thickness; WTI-M, intimal-medial layer thickness; ␭z, axial stretch ratio. *Statistically significant difference between the 2
groups (P ⬍ 0.05).
J Appl Physiol • VOL
Fig. 2. Representative example of a comparison of the incremental stressstrain relationships of a control and flow-overload RCA. Pressure was varied
in the 80- to 120-mmHg range and axial stretch ratio (␭z) was varied from 1.3
to 1.5.
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Pressure-overload in the right ventricle was induced by
stenosis of the PA for 4 wk which led to hypertrophy of the
right ventricle and hence an increase in RCA flow. Table 1
summarizes the characteristics of the experimental group relative to controls. The systolic RV pressure increased significantly after banding (159% of control value after 4 wk). The
increase in RV mass was also significant as reflected by the
138% increase in RV/(left ventricle ⫹ septum) ratio compared
with control hearts. The mean arterial blood pressure was not
significantly different in the two groups (64 ⫾ 6 vs. 68 ⫾ 8
mmHg for control and experimental groups, respectively).
The flow rate in the RCA increased by ⬃104% from a mean
control value of 39.9 ml/min. The inner diameter of the RCA
increased approximately by 15% which nearly normalized the
WSS to the control value of ⬃10 dyn/cm2. The pressure-outer
diameter relationship for the experimental and control animal
is shown in Fig. 1. It is clear that the remodeled vessel has a
larger outer diameter at any given pressure. This also reflects
the changes in the inner diameter since the vessel wall thickness (including intima-medial and adventitial layer) is not
significantly different in the two groups as shown in Table 1.
The inner diameter was calculated, using the incompressibility
assumption, from the outer diameter, the axial stretch ratio, and
the no-load area as per Eq. 2 in the APPENDIX. Figure 1 also
demonstrates a greater rate of change in diameter at lower
pressures for the control compared with the experimental
vessel which indicates loss of compliance in the experimental
vessel.
The circumferential strain and stress were computed according to Eqs. 1a and 3, respectively (APPENDIX), in reference to the
zero-stress state. The zero-stress state can be characterized by
the opening angle which was found to decrease from 156 ⫾
17.8 degrees in the control to 131 ⫾ 10.5 degrees in the
experimental group. The incremental circumferential stressstrain data are shown in Fig. 2 for various ␭z of a control and
experimental animal. The leftward shift and increase in slope
suggest stiffening of the vessels in the experimental group. The
axial strain and second Piola-Kirchhoff stress were also com-
STIFFER RCA IN FLOW-OVERLOAD
Fig. 3. Data (in kPa) of circumferential (Y11), axial (Y22), and cross (Y12 ⫽ Y21)
incremental elastic moduli for control (mean stress of 32–53 kPa) and experimental (mean stress of 42– 63 kPa) RCA. *Statistically significant difference
between the 2 groups (P ⬍ 0.05).
J Appl Physiol • VOL
and elastin (79%) for the wall. In the wall, collagen increased
significantly from an average value of 11.6 ⫾ 0.40 to 37.7 ⫾
2.8% (P ⬍ 0.001). Elastin also increased significantly in the
wall from 6.6 ⫾ 0.87 to 11.8 ⫾ 1.2% (P ⫽ 0.007) for the
control and experimental stacks, respectively.
Figure 6 summarizes the collagen-to-elastin (C/E) ratios of
the wall and individual layers (media and adventitia) of the
control and remodeled vessels. From Fig. 6, the C/E ratio for
the media increased significantly from a value of 4.0 ⫾ 0.53 to
6.6 ⫾ 0.38 (P ⫽ 0.005). In the adventitia, there was no
significant change in the C/E ratio in the experimental group
(1.1 ⫾ 0.10 vs. 1.3 ⫾ 0.12; P ⫽ 0.392). The C/E ratio of the
wall increased significantly from 1.8 ⫾ 0.23 to 3.3 ⫾ 0.33
(P ⫽ 0.008).
To further examine the changes in the arterial wall, and to
evaluate smooth muscle hyperplasia, histological sections were
prepared. Figure 7 demonstrates the arterial wall from a control
(Fig. 7A) to a remodeled (Fig. 7B) vessel. Elastic fibers are
visualized in blue-black to black, collagen is pink to red, nuclei
are blue to black, and other tissue elements are yellow to
brown. The number of nuclei per millimeter squared section
area did not change significantly from 616.7 ⫾ 56.5 to 569.9 ⫾
53.7 (P ⫽ 0.285) for the control and experimental segments,
respectively. The small and insignificant difference (8%) in
nuclei per area confirmed the absence of hyperplasia in this
study. The images in Fig. 7 represent between 15 to 20% of the
vessel ring circumference.
DISCUSSION
Previous studies established that the mechanical properties
of vessels are essentially determined by the matrix components
of the wall (49). Thus, changes in composition and structure of
the wall will result in altered vessel wall mechanics. Here, we
show that flow-overload induces an increase in the elastic
moduli in the axial and circumferential directions (increase in
stiffness). In the present model, the microstructural basis for
the increase in stiffness correlates with an increase in elastin
and collagen in the media and particularly in the adventitia
layer. This underscores, in the RCA, the importance of the
adventitia during expansive remodeling.
Flow-overload model. Our current model of RVH has the
effect of flow increase in the RCA (Table 1) with no change in
pressure (40). In this model, the banding or “obstruction” of the
Fig. 4. Means ⫾ SE of Cauchy stress (kPa) and Green strain at physiological
loading (pressure of 100 mmHg and axial stretch ratio of 1.4) for control and
experimental RCA groups. *Statistically significant difference between the 2
groups (P ⬍ 0.05).
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demonstrate raw data of an overlay (top left) of SHG/TPEF
along with individual signals (top middle: SHG; top right:
TPEF) obtained from the porcine RCA wall for a control (Fig.
5A) and remodeled (Fig. 5B) specimen. Through overlapping
of both SHG and TPEF, a high-contrast image that shows
structural details of the adventitia and media is attained. The
bottom rows of Fig. 5, A and B, illustrate black and white
representations of the top row (collagen and elastin) with the
brightness and contrast adjusted for optimum visualization.
These images correspond to ⬃8.3 to 12% of the wall circumference.
For a semiquantitative analysis, the imaged slices of the
entire wall thickness of coronary specimens were processed
with the ImageJ software available from the National Institutes
of Health. The elastin and collagen contents were analyzed for
their relative area fractions, calculated with respect to the total
area in the region of interest. Table 2 summarizes the average
area measurements (per two-dimensional image) of elastin and
collagen in the wall and individual layers (media and adventitia) of the control and experimental groups. Table 2 also
includes the percent change in elastin and collagen content in
the wall and individual layers (media and adventitia) for
complete stacks from control and experimental groups. Values
for the wall correspond to a region of 230 ⫻ 230 ␮m and
values for the individual layers (media and adventitia) correspond to localized 50 ⫻ 50-␮m regions of interest within the
wall. The values were normalized for stack size which ranged
between 42 and 115 images (21–57.5 ␮m).
Table 2 shows that collagen increased significantly in the
media and adventitia by ⬃91 and 112%, respectively. In the
media, the average percent content of collagen increased significantly from 19.8 ⫾ 0.86 to 37.9 ⫾ 3.0% (P ⬍ 0.001) for the
control and experimental stacks, respectively. In the adventitia,
collagen increased significantly from an average percent content value of 28.5 ⫾ 1.01 to 60.5 ⫾ 2.04% (P ⬍ 0.001).
Elastin, however, had no significant change in the media (9%
of control value) but increased significantly in the adventitia by
⬃80% (Table 2). In the media, the elastin did not change
significantly from an average percent content value of 5.3 ⫾
0.74 to 5.8 ⫾ 0.36% (P ⫽ 0.578) for the control and experimental stacks, respectively. In the adventitia, elastin increased
significantly from 25.7 ⫾ 2.4 to 46.4 ⫾ 3.0% (P ⬍ 0.001).
Table 2 also shows a significant increase in collagen (226%)
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Fig. 5. Two-dimensional visualization [2-photon excited fluorescence (TPEF)/second-harmonic generation (SHG)] of coronary arterial walls (cross sections
containing adventitial, medial, and intimal layers of RCA). The images are 230 ⫻ 230 ␮m. The excitation wavelength is ␭z ⫽ 800 nm. A: control segment.
B: experimental segment. Top, middle: collagen (from SHG). Top, right: elastin (from TPEF). Left: overlap of the signals. Bottom: illustrates black and white
representations of the collagen and elastin in the top. Magnification 40⫻.
J Appl Physiol • VOL
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STIFFER RCA IN FLOW-OVERLOAD
1343
Table 2. Average area measurements (per 2D image)
of elastin and collagen in RCA
Wall
Elastin
Collagen
Media
Elastin
Collagen
Adventitia
Elastin
Collagen
Control
RVH
%Change in Content
3,517⫾459
6,132⫾211
6,278⫾618*
19,994⫾1,485*
79%*
226%*
132⫾19
494⫾21
144⫾9
946⫾75*
9%
91%*
642⫾60
710⫾25
1,157⫾75*
1,508⫾51*
80%*
112%*
Values are means ⫾ SE (␮m2). Values for the wall correspond to a region
of interest of 230 ⫻ 230 ␮m and values for the individual layers (media and
adventitia) correspond to localized 50 ⫻ 50-␮m regions of interest. *Statistically significant difference between the 2 groups (P ⬍ 0.05). 2D, 2-dimensional.
Fig. 6. Summary of collagen-to-elastin (C/E) ratios of the RCA wall and
individual layers (media and adventitia) for control and remodeled (flowoverload) groups. *Statistically significant difference between the 2 groups
(P ⬍ 0.05).
J Appl Physiol • VOL
Fig. 7. Representative histological sections of RCA cross-sections. A: control
segment. B: experimental segment.
Remodeling of adventitia and microstructure. The use of
MPM for the visualization of elastin and collagen in the
remodeled wall is novel to this study. Previous studies employed tissue processing techniques to quantify only the content of collagen and elastin but not the arrangement of structure
(49). MPM allows for the visualization and quantification of
collagen and elastin (Fig. 5) without the need for staining,
processing, or digestion techniques. This new approach allows
for visualization of the three-dimensional structure which is
fundamental for future microstructural modeling. Although
previous studies utilized MPM to detect microscopic anatomical changes in vessels (49), this study is the first to utilize the
three-dimensional rendering capabilities to measure changes in
content (Table 2).
The individual wall components (mainly collagen and elastin fibers) are loaded sequentially which is the cause of the
nonlinear response in the stress-strain relationship (49). Elastin
(with a low elastic modulus) is preferentially the initial loadbearer, while collagen (with a high elastic modulus) is recruited as the load causes a high stretch. The variation of the
incremental elastic moduli, therefore, is said to be an indicator
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PA causes a pressure increase in the right ventricle which leads
to concentric hypertrophy, an adaptation of the ventricle to
reduce myocardial wall stress (18). Consequently, an increase
in RV mass develops which increases the blood supply demand
and therefore necessitates a gradual increase in flow in the
RCA as reported previously using microsphere blood flow
measurements (48).
Remodeling of RCA stiffness in RVH. Although the increase
in stiffness is a well-known phenomenon in the hypertensive
vessels (3, 7, 21), vascular enlargement in flow-overload has
previously been considered a physiological response to restore
the WSS. This is the first study to demonstrate increased
stiffness of a blood vessel in chronic flow-overload (Figs. 2 and
3). Here, we quantitatively characterized the change in mechanical properties (Fig. 3) for the intact artery ex vivo. The
literature on the mechanical properties of isolated remodeled
vessels is very limited (28). Those that do provide mechanical
data focus only on distension with no regard to the axial stretch
(28). Our study shows that differences in axial strain between
normal and remodeled vessels are significant (Figs. 3 and 4).
Other studies focused primarily on residual strain measurements of opening angles (10, 20). Although this is important, it
does not provide a thorough understanding of the mechanics of
the vessel wall. Our study is the first to quantify the three
incremental elastic moduli of a remodeled vessel in reference
to the zero-stress state (Fig. 3).
1344
STIFFER RCA IN FLOW-OVERLOAD
J Appl Physiol • VOL
implies that the vascular system closely regulates the degree of
deformation. The regulation of strain was clearly observed in
the animal model after 4 wk of flow-overload (Fig. 4).
Remodeling of the zero-stress state. Fung (9) previously
proposed that the remodeling of the zero-stress state is an index
of the nonuniformity of growth and remodeling. Fung and Liu
(10, 12, 35) showed that hypertension induces growth of intima
that exceeds that of the adventitia. Consequently, the vessel
sector (in the zero-stress state) shows an outward bend and
hence an increase in the opening angle. Conversely, Lu et al.
(39) showed that flow-overload induces growth of adventitia
that exceeds that of intima which is consistent with our current
findings in the flow-overload model of RCA (Table 2). Hence,
the vessel sector bends inward and decreases the opening angle
as reported in this study. In summary, hypertension and flowoverload have opposing effects on the opening angle. In
models of simultaneous hypertension and flow-overload, the
effects of opening angle were found to be conciliatory (25).
Comparison with other studies. Our group recently determined the values of the elastic moduli of a normal intact RCA
and its layers (intima-media and adventitia) in normal pigs
(37). Although the results are generally similar for a control
vessel, a methodological difference should be noted. A major
difference between our previous and present measurements is
the implementation of ultrasonic crystals for the measurement
of dimensions during loading (see Fig. 1). This allows for
continuous data acquisition during testing of a vessel and hence
reveals primarily the elastic response. This differed from previous studies where a camera was used to take snapshots to
record dimensions at discrete points in which relaxation of
stress as well as creep may have occurred. To our knowledge,
there are no equivalent data on the coronary arteries in flowoverload.
Critique of methods. The extent of manipulation of the RCA
for placement of a flow transducer warrants a comment. The
site of placement of the transducer was proximal to the vessel
emergence from the aorta and the site of testing of measurement was 3–3.5 cm away. Hence, the vessel of interest was
sufficiently far away from area of injury. We compared vessels
in models with and without the placement of a flow transducer
and found no significant differences. Histological examination
verified the lack of injury in the vessel wall (Fig. 7).
Stress-strain behavior and the incremental elastic moduli
were implemented to quantitatively describe the stiffness of the
isolated RCA. A limitation here is that we focused on a
relatively small pressure range (80 –120 mmHg) to linearize
the stress-strain curves. Although this is sufficient to determine
the in vivo properties, a full range of deformation analysis
would allow a more thorough understanding of the biomechanical behavior of these vessels. Finally, we considered the
vessel wall as a homogenous structure. Given that the incremental moduli are layer (intima-media and adventitia) dependent, the methods employed in this study should be applied to
the individual layers to investigate their role and contribution
to the remodeling of the coronary artery.
Significance of study. Stress and strain are intimately related
to tissue function, growth, and remodeling. Hence, a thorough
understanding of the stress and strain state in normal and
diseased vessel wall can be used to predict vascular dysfunction. Here, we observed changes in vascular stiffness that
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of vessel structural composition, and is observed to change
significantly with vascular remodeling. In this study, the collagen content increased significantly both in the media and
adventitia (Table 2). Elastin, on the other hand, had almost no
change in the media, although there was a significant increase
in the adventitia. The increase in microstructural content did
not significantly affect the vessel wall thickness (Table 1),
suggesting a nonuniform growth. A prevailing growth in the
adventitial layer concurrently with an unchanging wall thickness is consistent with previous flow-overload studies (39).
From Fig. 6, the results show an increase in the C/E ratios of
⬃77.5% overall (wall), 63% in the media, but only 12% in the
adventitia from control to experimental vessels. These data
suggest that RVH stress stimulates the induction of collagen
more rapidly than elastin in the media but at a similar rate
in the adventitia. The significant increase in collagen and
elastin, particularly in the adventitial layer, may explain the
increase in stiffness in our model. This shift in microstructural
composition may adversely affect or limit the vessel’s functional abilities. Thus, the previously thought physiological
compensatory enlargement in flow-overload (Fig. 1) may lead
to vascular dysfunction.
Uniform shear hypothesis. The remodeling of blood vessels
in response to flow-overload obeys the constant WSS hypothesis (23, 39). This principle implies that the volumetric flow
rate is proportional to the cube of the vessel radius assuming a
laminar, incompressible, Newtonian flow through a rigid cylindrical vessel (24). Hence, to maintain a constant WSS, the
cube of the radius must increase in proportion to the increase
in blood flow. Blood vessels accommodate such a change in
vessel radius acutely through vasoactive regulation, and chronically by remodeling vascular caliber, and thus maintaining
homeostasis (25). The RCA was found to undergo enlargement
of the caliber (Fig. 1) to restore the WSS (Table 1) which is
consistent with previous studies in other vessels (2, 25, 30, 32).
Uniform circumferential stress and strain hypothesis. An
increase in mean circumferential wall stress (Fig. 4) was
expected, according to APPENDIX Eq. 3a, because the vessel
dilates (increases in radius) in flow-overload with no significant change in wall thickness (see Table 1). This is consistent
with Girerd et al. (15) who reported a 51% increase in circumferential stress despite normalization of WSS in an arterialvenous (a-v) fistula of patients with end-stage renal disease
(ESRD). The strain, on the other hand, normalizes to the
control values fully after 4 wk (Fig. 4). These findings are in
agreement with our previous longitudinal hypertension and
flow-overload studies (5, 39). In both models, strain reached its
peak sooner and normalized faster than mean wall stress.
Hence, we concluded that the vessel appears more “sensitive”
to changes in strain which implies the enforcement of a
homeostatic state of strain.
In support of the notion of strain homeostasis in the vascular
system, Guo and Kassab (17) recently determined the distribution of circumferential stress and strain along the porcine
aorta and throughout the coronary arterial tree. They showed
that the circumferential stretch ratio and average wall stress
varied from 1.2 to 1.6 and 10 to 150 kPa, respectively, along
the aorta and the entire left anterior descending coronary
arterial tree (more than 3 orders of magnitude difference in
vessel diameters). The relative uniformity of circumferential
stretch ratio from the proximal aorta to a 10-␮m arteriole
1345
STIFFER RCA IN FLOW-OVERLOAD
APPENDIX
Theoretical Formulation
The circumferential deformation of an artery may be described by
the midwall circumferential Lagrangian-Green strain, which is defined
as follows
1
E ␪ ⫽ 共␭ ␪2 ⫺ 1兲
2
(1a)
where ␭␪ is the midwall stretch ratio (␭␪ ⫽ c/C, where c refers to
the midwall circumference of the vessel in the loaded state and C
refers to the corresponding midwall circumference in the zero-stress
state). Similarly, the axial Green strain is given by
1
E z ⫽ 共␭2z ⫺ 1兲
2
(1b)
where ␭z is the axial stretch ratio (defined as the change in axial
length between loaded and no-load state).
The midwall circumference in the loaded state was computed from
the average of inner and outer diameter. The inner diameter (Di) of the
vessel can be computed from the incompressibility condition for a
cylindrical vessel as
冑
D i ⫽ Do2 ⫺
4Ao
␲␭z
(2)
where Do and A0 are outer diameter at the loaded state and wall
area in the no-load state, respectively. Because all the quantities on the
right-hand side of Eq. 2 are measured, the loaded Di can be computed.
The total wall thickness (h) was computed as h ⫽ (Do ⫺ Di)/2.
The mean Cauchy stress in the circumferential (␶␪) is given by
Laplace’s equation as
␶␪ ⫽
PD i
2h
(3a)
where P is the luminal pressure. The second Piola-Kirchhoff
circumferential stress (S␪) is related to the Cauchy stress by
␶␪
S␪ ⫽ 2
␭o
Similarly, the axial second Piola-Kirchhoff stress (Sz) is given by
Sz ⫽
冋
册
4F
1
PD2i
2
2
2 ⫹
␭z ␲共Do ⫺ Di 兲 2h共Do ⫹ Di兲
(4)
where F is the axial force imposed on the vessel.
Incremental Moduli
If we assume the material of the blood vessel wall is homogeneous,
incompressible, orthotopic, and linearly elastic, the classical theory of
thin-walled elastic shells can be used to compute the linearized
stress-strain relationships and hence the elastic moduli. If a small
perturbation of stress and strain from a homeostatic in vivo state is
considered, a linear incremental stress-strain relationship can be
expressed as follows
␦S 11 ⫽ Y 11 ␦E 11 ⫹ Y 12 ␦E 22 ,
␦S 22 ⫽ Y 21 ␦E 11 ⫹ Y 22 ␦E 22
(5)
where Sij and Eij are components of second Piola-Kirchhoff stress
and Lagrange-Green strain, respectively. The indexes 1 and 2 denote
circumferential and axial directions, respectively. Y11 and Y22 are the
classic incremental Young’s moduli in the circumferential and axial
directions, respectively; Y12 ⫽ Y21 are denoted as cross-modulus,
although they have no equivalents in classic mechanics but are
necessary terms for the strain energy function (11). Although these
equations are Hookean, they are not isotropic.
Least-Squares Method for Determination of Elastic Moduli
Three elastic moduli (i.e., Y11, Y12, and Y22) must be determined as
given by Eq. 5. The method of derivation of elastic moduli is clearly
outlined by Huang et al. (20). Briefly, a least-squares method is used
to minimize the error between the theoretical stresses given by Eq. 5
and the experimental measurements. The result is a 3 ⫻ 3 matrix
whose solution is the 3 ⫻ 1 matrix of elastic moduli
冋
A 11
0
A 12
0
A 22
A 21
A 12
A 21
A 11 ⫹ A 22
册冋 册 冋 册
Y 11
B 11
Y 22 ⫽ B 22
Y 12
B 12
(6)
where
n
n
A 11 ⫽ ⌺ n ⌬E 11
⌬E 11
n
n
A 12 ⫽ A 21 ⫽ ⌺ n ⌬E 11
⌬E 22
n
n
A 22 ⫽ ⌺ n ⌬E 11 ⌬E 22
n
n
B 11 ⫽ ⌺ n ⌬S 11
⌬E 11
n
n
n
n
B 12 ⫽ ⌺ n ⌬S 22
⌬E 11
⫹ ⌺ n ⌬S 11
⌬E 22
n
n
B 22 ⫽ ⌺ n ⌬S 22 ⌬S 22
and
n
n
0
⌬S 11
⫽ S 11
⫺ S 11
,
n
n
0
⌬S 22
⫽ S 22
⫺ S 22
,
n
n
0
⌬E 11
⫽ E 11
⫺ E 11
,
n
n
0
⌬E 22
⫽ E 22
⫺ E 22
,
n ⫽ 0,1,2 . . .
The in vivo state is denoted by n ⫽ 0 and consequently
0
0
0
0
⫽ ⌬S 22
⫽ ⌬E 11
⫽ ⌬E 22
⫽0
⌬S 11
The symbol ␦ is dropped for convenience, but it should be recalled
that the quantities of stress and strain are all defined in the incremental
sense according to Eq. 5. The constants A and B are determined from
the n experiments, and the matrix is solved for the three elastic
moduli.
ACKNOWLEDGMENTS
We thank Dr. Q. Dang for excellent technical expertise.
GRANTS
(3b)
J Appl Physiol • VOL
This research was supported in part by National Heart, Lung, and Blood
Institute Grants HL-087235 and HL-084529.
106 • APRIL 2009 •
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accompanies the expansive remodeling of coronary arteries in
flow-overload. Given that increased stiffness is an important
risk factor for vascular dysfunction, flow-overload in RVH
may predispose the RCA to endothelial dysfunction and vascular disease. In addition to cardiac hypertrophy, there are
many possible clinical conditions in which the findings of this
study may have an impact. Other models of flow-overload
include chronic obstructive pulmonary diseases, Bernheim’s
Syndrome, mitral valve stenosis, and a-v fistula used in dialysis
for ESRD patients. These conditions may involve other factors
which contribute to their pathology in addition to flow-overload. For instance, in patients with ESRD, the role of uremiaspecific factors (mainly calcification) may significantly affect
the observed arterial stiffness (29). The influences of calcification, enzymes, cytokines, and other proteins have yet to be
completely understood in these conditions. In our study, the
model allowed for the determination of the effect of pure
flow-overload on the RCA. The synergistic interplay of flowoverload with biochemical factors in other disease states remains to be investigated in future studies.
1346
STIFFER RCA IN FLOW-OVERLOAD
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