changes in the structure-function relationship of elastin and its impact on the proximal pulmonary...

Changes in the structure-function relationship of elastin and its impact on the proximal pulmonary arterial mechanics of hypertensive calves Running Head: Elastin mechanics in pulmonary hypertension

Steve Lammers1, Phil Kao1, H. Jerry Qi1, Kendall Hunter2, Craig Lanning2, Joseph Albietz3, Stephen Hofmeister3, Robert Mecham4, Kurt R. Stenmark3, Robin Shandas1,2

(1) Mechanical Engineering

University of Colorado

Boulder, CO

(2) Dept. of Pediatric Cardiology

University of Colorado Health Sciences

Denver, CO

(3) Department of Pediatrics

Developmental Lung Biology Laboratory

University of Colorado Health Sciences

Denver, Colorado

(4) Dept. of Biology and Biomedical Sciences

Washington University

St. Louis, MO

Articles in PresS. Am J Physiol Heart Circ Physiol (July 25, 2008). doi:10.1152/ajpheart.00127.2008

Copyright © 2008 by the American Physiological Society.

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Address for Correspondence:

Robin Shandas, Ph.D. Professor Center for Bioengineering University of Colorado 13123 E. 16th Avenue, B100 Aurora, CO 80045 Robin.shandas@uchsc.edu Phone: (720) 777 2586 Fax: (720) 777 4056

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Abstract

Extracellular matrix remodeling has been proposed as one mechanism by which proximal

pulmonary arteries stiffen during pulmonary arterial hypertension (PAH). Although some

attention has been paid to the role of collagen and metallomatrix proteins in affecting vascular

stiffness, much less work has been performed on changes in elastin structure-function

relationships in PAH. Such work is warranted given the importance of elastin as the structural

protein primarily responsible for the passive elastic behavior of these conduit arteries. Here, we

study structure-function relationships of fresh arterial tissue and purified arterial elastin from the

main, left and right pulmonary artery branches of normotensive and hypoxia-induced pulmonary

hypertensive neonatal calves. PAH resulted in an average 81% and 72% increase in stiffness of

fresh and digested tissue respectively. Increase in stiffness appears most attributable to elevated

elastic modulus, which increased 46% and 65% respectively for fresh and digested tissue.

Comparison between fresh and digested tissues shows that, at 35% strain, a minimum of 48% of

the arterial load is carried by elastin, and a minimum of 43% of the change in stiffness of arterial

tissue is due to the change in elastin stiffness. Analysis of the stress-strain behavior revealed that

PAH causes an increase in the strains associated with the physiologic pressure range but had no

effect on the strain of transition from elastin-dominant to collagen-dominant behavior. These

results indicate that mechanobiological adaptations of the continuum and geometric properties of

elastin, in response to PAH, significantly elevate the circumferential stiffness of proximal

pulmonary arterial tissue.

Key Words: pulmonary hypertension, elastic modulus, stiffness, vascular remodeling

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Introduction

Pulmonary arterial hypertension (PAH) in neonates, infants and children is a progressive

condition marked by substantial vascular remodeling of the distal pulmonary vasculature, which

increases pulmonary vascular resistance (PVR) and pulmonary arterial pressure. PVR

measurement has been the standard diagnostic for evaluating both the significance of PAH and

the extent of vascular response to therapies. However, PVR is an inherently limited diagnostic in

that it ignores pulmonary vascular stiffness (PVS), an especially important omission given the

inherently pulsatile nature of cardiac function and the importance of robust ventriculo-vascular

coupling in maintaining hemodynamic efficiency through the pulmonary vasculature. In fact,

many studies of vascular function in systemic hypertension are documenting the substantial role

played by the elastic proximal arteries in maintaining systemic vascular hemodynamic efficiency

and reducing cardiac workload. Several investigators have shown the significant mechanical

advantages conveyed by the elasticity of systemic conduit arteries in reducing overall hydraulic

impedance and cardiac workload (29, 32, 37, 38). Other authors have correlated proximal artery

stiffness and cardiovascular mortality for patients with systemic hypertension (5, 13, 28).

Although less work has been done regarding pulmonary artery (PA) elasticity in PAH, several

recent animal studies are beginning to address this topic. In particular, recent work by Kobs and

Chesler used fresh artery material testing to correlate increased collagen and elastin content with

elevated elastic moduli for the collagen and elastin dominant regions of the stress-strain curve

respectively(26).

In addition, several clinical studies have shown that increased PVS is correlated with

patient mortality (12, 18, 31, 43) and can account for as much as 30% to 40% of the increased

load on the heart (40) due to elevated vascular hydraulic impedance and right ventricular

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afterload. Remodeling of the extra-cellular matrix (ECM) in particular appears to be an

important aspect of PVS and a number of studies have examined various biological mechanisms

including inflammatory factors, upregulation of proliferative catalysts, and phenotypic changes

in resident cells (2-4, 6, 17, 20-24, 40, 41, 44). Other studies have shown increased collagen

content and vascular tone in stiffer arteries (7, 11, 14, 19, 21, 27, 34, 35). In elastin specific

studies, elastolytic activity has been shown to increase during PAH (33), and inhibition of this

pathway by the serine elastase inhibitor elafin mitigates hypoxia-induced PAH (46). In recent

work from our group coupling a novel orthotropic, hyperelastic, microstructural model of the PA

wall to biomechanical studies of PAs from normotensive and hypertensive rats, we suggested

that increased ECM structural protein crosslinking may be one mechanism by which these

vessels stiffen (47, 48). The next step in addressing this hypothesis is to evaluate

experimentally whether the increased stiffness seen in hypoxia-induced PAH can be attributable

to increased stiffness of the structural protein matrix. Given that elastin is the ECM protein

largely responsible for vascular elasticity, we evaluate here in a well-established PAH neonatal

calf model the role of elastin in modulating the stiffness of upstream PA vessels.

In this study, we hypothesize that changes in the functional mechanical properties of

elastin is one of the primary causes of the increased stiffness seen in the proximal PAs due to

hypoxia-induced PAH. To test this hypothesis, we examined the biomechanics and structure of

normotensive and hypertensive neonatal calf proximal PAs and their associated elastin tissue.

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Materials and methods

Animal model

All studies were performed after approval by institutional animal care and use

committees. Arterial tissues from 3 normotensive and 6 hypertensive male Holstein calves (70-

110lb) were used. Hypertension was induced by hypobaric hypoxia, 2wks, 430 mmHg [4600 m

equivalent air pressure]. Hypoxic animals were sacrificed at 430 mmHg pressure after 14 days at

hypoxic conditions. All animals were studied and sacrificed at 15 ± 1-2 days of age. PA

pressures were obtained with a 5-French Swan-Ganz or Millar catheter, positioned in the main

PA. After sacrifice the right, left and main PAs (RPA, LPA, MPA) were excised, stored in

nutrient balanced medium (4 ºC). Testing was conducted within 24 hours of sacrifice, material

properties of fresh tissues were unchanged for tests conducted up to 4 days post sacrifice.

Geometric and material property results were calculated using 3 control and 6

hypertensive tissues for the LPA and MPA. Two hypertensive RPA tissues were too small and

irregularly shaped for stress-strain testing; therefore 3 control and 4 hypertensive RPA tissues

were used. We were also unable to accurately measure the pulmonary arterial pressure for one

hypertensive animal. Due to this, 3 control and 5 hypertensive LPA and MPA and 3 control and

4 hypertensive RPA tissues were used for all material property comparisons made at physiologic

pressure/strain values.

Tissue dissection and processing

Arteries were inspected for holes, tears, branch points and localized thickening or other

anomalies that may adversely affect test results. Loose connective tissue was removed, and

gross anatomical measurements of diameter (1-distal, 1-proximal) and wall thickness (4-distal, 4-

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proximal) were obtained with digital calipers prior to axial bisection, see figure 1.

Circumferential rectangular sections (~8mm x 20mm) were taken for subsequent material

testing. Biopsy samples were taken for histological analysis, three each in the longitudinal and

circumferential directions.

Materials testing

Fresh tissue stress-strain data was acquired with an MTS, Insight 2 (MTS Systems

Corporation, Eden Prairie, MN) material testing system, equipped with an ancillary

environmental chamber and either 5N or 2kN load cell, depending on the maximum expected

load for each test, figure 2. Stress-strain tests were conducted in Calcium and Magnesium-free

NaCl-PBS buffer (.01 mol/L, ionic str. 0.15, pH 7.4) 36ºC. Sample width, thickness and gage

length were measured using digital calipers. Stress-strain properties remained consistent for

tissue tests performed up to three days post sacrifice.

Tissues were preconditioned by ten full extension-relaxation cycles before data were

taken. The strain rate for all mechanical tests was 10% strain per second. In this study, force

and stress data refer to the loading region of the stress-strain curve. In an effort to avoid damage

that may affect elastin-only material tests, tissues were stretched until the collagen-dependent

strain-stiffening response became fully developed, typically 60 – 80% strain, figure 3. Elastin

tissue tests were conducted in the same manner as that outlined for fresh arteries, but with a 2N

or 5N load cell. Elastin samples were tested at strains ranging from 60 to 100% at 10%

increments or until failure. Data processing was conducted using custom-written software

(Matlab® R2006a, The Math Works, Inc., Natick, MA).

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Tissue digestion using formic acid – CNBr solution

After fresh artery material testing, tissues were processed into elastin-only scaffolds using

the CNBr-formic acid-digestion procedure outlined by Lu et al (30). Briefly, fresh arteries were

submerged in a 70% formic acid, 50 mg/mL Cyanogen Bromide (CNBr) solution and gently

stirred for 19 hours at room temperature followed by 1 hour at 60ºC. The samples were then

boiled for 10 min. to remove any residual CNBr. The digestion solution was then decanted and

the tissue was rinsed with distilled water.

CNBr-formic acid-digestion was chosen for elastin purification due to its effectiveness

for removing non-elastin components from arterial tissues. Other common elastin purification

procedures using hot alkali or autoclaving were rejected due to possible network fragmentation

and incomplete elastin purification respectively (15).

Amino acid analysis was performed on LPA elastin samples (n=7) to determine the

effectiveness of the digestion procedure. After digestion, the samples retain formic acid and

subsequent material testing showed strong pH dependence. NaOH titration was therefore used to

raise the pH from 2-3 to 7.4. After titration, samples were stored in NaCl-PBS buffer and were

tested within 2 weeks of animal sacrifice. Elastin mechanical properties were unchanged for

tests conducted up to one month post sacrifice.

Histology and elastin area fraction

Biopsies were fixed in 10% buffered formalin, wax embedded, sectioned and stained for

elastin using Verhoeff’s Van Gieson. Bright-field photomicrographs were taken with a Nikon

TE-200 (Nikon Instruments Inc., Melville, NY) microscope (40-X magnification, SPOT RT-900

camera [Diagnostic Instruments Inc., Sterling Heights, MI]). Elastin area fraction was

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determined using custom software (Matlab® R2006a Image Processing Toolbox, The Math

Works, Inc., Natick, MA).

Calculations

Fresh artery Cauchy stress (σfresh) was calculated with the assumption of

incompressibility applied to the width (W) and thickness (T) directions to account for changes in

cross-sectional area as a function of engineering strain (ε).

WTF

fresh =σ elast

elast AWTF0=σ (1)

( ) νε −+= 10TT ( ) νε −+= 10WW (2)

0

0

LLL −=ε (3)

Where F is the applied force and L is the loaded sample length in the direction of applied load.

T0, W0 and L0 are the initial thickness, width and length respectively, Aelast is the area fraction of

elastin and υ is Poisson’s ratio (0.5 for incompressible tissue). Due to the layered structure of the

fenestrated elastic sheets, stresses for the digested elastin samples (σelast) were calculated with the

condition of incompressibility applied to the width direction only. To facilitate modulus (Ε)

calculations, stress-strain data for fresh and elastin samples were fit with 9th and 4th order

polynomials respectively. The 9th and 4th order polynomials were determined to be the lowest

order polynomials necessary to achieve an average squared correlation coefficient R2avg >=0.995

for all fresh and elastin datasets respectively.

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εσ

dd=Ε (5)

Stiffness (Φ) is the derivative of the width-normalized force (Fn: force divided by current

width) by strain, and as such, stiffness is dependent upon thickness geometry and represents the

extensive equivalent of elastic modulus.

( ) ( )( )εεεε nF

dd

WF

dd =⎟⎟

⎠

⎞⎜⎜⎝

⎛=Φ (6)

The condition of incompressibility was applied to the stiffness calculation to account for any

variation in width as a function of imposed strain. Fn data were fit to polynomials in the same

manner as was used for the stress. Unless otherwise noted, all results for the material parameters

(Ε, Φ) were calculated at a strain of 35%, so that tissues could be compared. Strains higher than

35% resulted in one or more of the stress-strain curves extending into the collagen-dominant,

strain-stiffening region. At strains greater then 35%, any comparison between fresh and elastin

stress-strain data would be complicated by the active loading of collagen. Since this collagen

component cannot be removed from the fresh artery dataset, any comparison of the mechanical

properties made at strains greater then 35% would include collagen and therefore not allow for

the direct comparison of elastin mechanical property changes from the fresh to digested state.

Lame’s equation for stress in thick-walled tubes (7) was used to determine the strains that

correspond to the measured physiologic pressures.

( )( )( ) 22

22

rTrTrrPi

L −+++

=σ , ( )ε+= 10rr (7)

Lame’s equation was equated to the polynomial stress-strain function and solved iteratively for

the strains that correspond to the in vivo measured systolic, diastolic and mean pressures (Pi); r

and r0 are the internal radii of the artery in the loaded and initial state respectively. The

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estimated physiologic strains were then used to calculate the systolic and diastolic moduli using

equation 5.

Curvature (κ) of the stress-strain curve was used to determine the onset of collagen

engagement, figure 4.

( ) 2321 σ

σκ′+

′′= (8)

The 2nd order derivative in the numerator of the curvature equation led to large variances when

calculated from the polynomial fit to the original stress-strain data. We therefore applied a zero-

phase, low-pass, elliptic digital filter to smooth the data. A new 9th order polynomial was then fit

to the filtered data, resulting in reduced curvature variance.

Average pre-strain-stiffening curvature (κPSS) was calculated for the strain range (εPSS)

between εL and εH, where εL equals 20% strain and εH equals the strain associated with maximum

curvature minus 20% strain, figure 4. εL and εH were defined in this way so that any curvature

associated with the low-strain loading region or high-strain strain-stiffening region would not

influence the κPSS value. Transition strain (εtrans) was then determined by the relationship:

100H

trans =κ PSSH κκ −= max (9)

such that εtrans is associated with the onset of curvature elevation. εtrans results compare

favorably to those previously published using both experimental(36) as well as Lorentz

distribution functions(45, 49). Stress-strain plots of elastin tissue do not exhibit curvature

elevation behavior, and since strain-stiffening is associated with collagen engagement, we

defined εtrans as the strain at which collagen begins to carry load. Strains below this value fall

within the elastin-dominant region of the stress-strain curve, A in figure 4, while strains which

are greater fall within the transitional or collagen-dominant region, B. We further defined the

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collagen dominant region (εcd) as those strains which result in a fresh artery modulus greater than

ten times the artery modulus calculated within the elastin dominant region at a strain of 35%.

While this is an arbitrary definition for collagen dominance it seemed logical to suggest that

collagen is the dominant load bearing structural protein when at least 90% of the material

modulus can be attributed to collagen.

Equation (10) defines the percent of the fresh tissue load that is carried by elastin (%Fn-C,

%Fn-H) and the percent of the increased load, due to PAH, which is attributable to the increased

stiffness of elastin (%Fn-e,f), calculated at a strain of 35%.

⎟⎟⎠

⎞⎜⎜⎝

⎛=

−

−−

Cfn

CenCn F

FF

,

,%% ⎟⎟⎠

⎞⎜⎜⎝

⎛=

−

−−

Hfn

HenHn F

FF

,

,%% ⎟⎟⎠

⎞⎜⎜⎝

⎛

−−

=−−

−−−

CfnHfn

CenHenfen FF

FFF

,,

,,, %%

Fn-f,C , Fn-f,H : Fn, fresh tissue (control, hypertensive).

Fn-e,C , Fn-e,H : Fn, elastin tissue (control , hypertensive).

(10)

Fn-C and Fn-H were similarly calculated at systolic and diastolic strains to determine the percent

load carrying capacity of elastin at physiologic strains.

P-values refer to the probability associated with the one-tail, two-sample equal variance,

Student’s t-test. Results were deemed statistically significant for P-values<=0.05. All error bars

represent one standard deviation above and below the mean. Standard deviations (S) were

calculated using the “n-1” method. In all curve fitting, the average of the squared correlation

coefficient was deemed statistically significant for R2avg >=0.995.

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Results

Morphology

PAH resulted in an increase in vessel wall thickness by 30% (S=12%, P=0.02), 28%

(S=11%, P=0.04) and 26% (S=8%, P=0.03,) for the RPA, LPA and MPA respectively, figure 5.

There was no statistically significant change in the internal diameter as a result of hypertension.

PAH also caused a decrease in the area fraction of elastin by 17% (S=2%, P=0.03), 16% (S=2%,

P=0.01) and 16% (S=3%, P=0.03) for the RPA, LPA and MPA respectively, figure 5. The

elastin content factor (ECF) is the product of the elastin area fraction and tissue thickness and is

representative of the volume of elastin in a given tissue cross section. The increase in ECF,

resulting from PAH, was calculated to be 7.6% (S=3.1%), 7.7% (S=3.1%) and 5.5% (S=2.0%)

for the RPA, LPA and MPA respectively, average ECF=6.9% (S=1.2%).

Physiology

Average systolic/diastolic pressures were 45/18 mmHg (S=12/7 mmHg) and 146/71 mmHg

(S=29/18 mmHg) for control and hypoxic populations respectively. Control and hypoxic mean

pulmonary pressures were 31 mmHg (S=12 mmHg) and 106 mmHg (S=16 mmHg) respectively.

We were unable to collect accurate diameter data, for all tested tissues; therefore, the strain at

systole and diastole were calculated using Lame’s equation (7) with average unloaded internal

radii calculated for each MPA, RPA and LPA; measured from 4 control and 11 hypertensive

tissues. The average systolic/diastolic strains for the PAs were 59% (S=10%) / 34% (S=6%)

and 72% (S=13%) / 53% (S=11%) for the control and hypoxic populations respectively. The

average strain for the mean pulmonary pressures were 47% (S=8%) and 65% (S=12%) for the

control and hypoxic populations respectively. The average transition strain was unchanged by

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PAH, 49% (S=8%) for control and 51% (S=9%) for hypertensive tissues (P=0.29). Transition

strain values are comparable to the strain corresponding with the onset of collagen engagement

published by Zulliger (49). The strain associated with the onset of collagen-dominance was

calculated to be 69% (S=7%) and 74% (S=8%) for the control and hypoxic populations

respectively.

Mechanical properties of fresh tissue

At 35% strain,the average increase in stiffness, due to PAH, was 72% (S=30%, P=0.01),

107% (S=56%, P=0.01) and 63% (S=14%, P=0.01) for the RPA, LPA and MPA respectively,

figure 6. The overall average increase in stiffness was 81% (S=23%, P<<0.05). PAH also

caused an in the increase in modulus of PA tissues by 28% (S=7%, P=0.04), 45% (S=23%,

P=0.10) and 65% (S=18%, P=0.02) for the RPA, LPA and MPA respectively. The average

overall increase in fresh tissue modulus was 46% (S=18%, P=0.02). Average stiffness values at

physiologic strains are shown in figure 7, and comparisons between physiologic stiffness values

are detailed in table 1.

Mechanical properties of purified arterial elastin

At 35% strain PAH resulted in an increase in the stiffness of the elastin by 57% (S=19%,

P=0.04), 109% (S=68%, P=0.03) and 50% (S=17%, P=0.05) for the RPA, LPA and MPA

respectively, figure 8. The overall average increase in the stiffness of elastin was 72% (S=33%,

P=0.01). The modulus of elastin increased by 34% (S=7%, P=0.04), 88% (S=43%, P=0.01) and

73% (S=36%, P=0.07) for the RPA, LPA and MPA respectively. The overall average increase in

the modulus of elastin was 65% (S=28%, P<<0.05). Elastin stiffness at physiologic pressure is

detailed in figure 7 and table 1. Amino acid analysis confirmed that CNBr-formic acid digestion

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resulted in elastin tissue samples of high purity(25), as evidenced by the high number of residues

per 1000 for alanine (227, S=20), valine (127, S=6) and glycine (342, S=16), along with the lack

of methionine and hydroxylysine residues (trace) and the low concentration of hydroxyproline

(14, S=5).

Average values for %Fn-C and %Fn-H are 51% (S=16%, P=0.02) and 48% (S=13%,

P=0.003) respectively, meaning that at a minimum, roughly half of the load applied to the fresh

arteries is carried by elastin for both control and hypertensive conditions at a strain of 35%.

Also, %Fn-e,f indicates that the change in the stiffness of the elastin accounts for an average of

43% (S=10%, P=0.003) of the change in the stiffness of the fresh tissue as a result of

hypertension. Changes in elastin load carrying capacity at physiologic strains are detailed in

table 1.

Discussion

Several key issues arise from this study. First, both the physiologic pressure and PA

stiffness elevation for the hypoxia-induced neonatal calf model mimic the changes we have seen

in the human (12), and confirm the value of this model for studying PA stiffness and impedance

changes in PAH. Second, the increases in stiffness appear attributable to a significant extent on

material rather than purely geometric factors. Third, the changes in material modulus due to

PAH appear to be significantly attributable to changes in elastin modulus specifically. Fourth,

the physiologic tissue strain at diastolic pressure is largely dependent on the stiffness of elastin.

Lastly, a significant component of the mechanical load for both normal and hypertensive

conditions continues to be carried by elastin. We comment on each of these issues below.

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Use of the Neonatal Bovine Model

The normal and pathologic physiology of the bovine model more closely mimics that of

humans than that of the available rodent models. Humans and young calves both have similar

body mass, heart rate, blood pressure and vascular wall thickness. In addition, the cellular

composition of the proximal PA wall of the large mammals is similarly complex to humans,

incorporating numerous phenotypically distinct smooth muscle cell populations which rapidly

proliferate during PAH (40, 41). For these reasons and the fact that we have significant

experience with this model, we chose the hypoxic neonatal calf as the in vivo model for

understanding how the proximal PA stiffens during hypoxia-induced PAH.

Impact of PAH on physiologic strain

Hypoxic conditions successfully induced PAH as evidenced by the significant increase in

mean PA pressure, vascular wall thickness and PVS. Due to the stress-strain nonlinearity of

arterial tissue, the relative contribution of elastin and collagen is contingent upon the strains

associated with in vivo physiologic pressures. Our results indicate that the physiologic region

of the stress-strain curve bounds the transition strain, with 62% of the curve residing in the

elastin-dominant region and 38% in the transition region for the control population, figure 9.

Hypertension shifts the physiologic strain region to higher strain values, but has no effect on the

transition strain. The resulting diastolic strain of hypertensive tissues is roughly equal to the

transition strain, 53% and 50% respectively, and the systolic strain of 72% nearly equals the

strain associated with collagen-dominant stress-strain behavior (74% in PAH

tissues).Hypertensive stress-strain curves therefore begin at the transition strain and reside

entirely within the transition region. This bounding of the transition strain in control tissues and

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the correlation of transition and diastolic strains for hypertensive tissues suggests that the

diastolic strain, and therefore baseline slope, of the stress-strain curve is set by the stiffness of

elastin, while the systolic strain incorporates collagen engagement.

At diastolic strain, PAH causes a 293% (ΦDias, Cont to PAH) increase in artery stiffness,

figure 7, table 1. Since diastolic strain resides within the elastin dominant region for control

tissues and at the transition strain for PAH tissues, this increase in diastolic stiffness is likely the

result of both increased elastin stiffness and the onset of collagen engagement. We calculate that

PAH causes a 173% (ΦDias, Cont to PAH) increase in diastolic elastin stiffness for digested samples.

The discrepancy in diastolic stiffness elevation between elastin and fresh tissue is a consequence

of the removal of non-elastin material during elastin purification, which will necessarily change

the mechanical properties of elastin between its natural and tested state. However, given that the

minimum percent diastolic load-carrying capacity of elastin is 61% (%Fn Dias, Cont) and 48% (%Fn

Dias, PAH) for the control and PAH tissues respectively, and that elastin stiffness values are

relatively strain independent, we feel confident in stating that the change in diastolic stiffness,

due to PAH, is predominately a result of an increase in elastin stiffness.

The significant strain stiffening response of the transitional region acts to limit the

maximum systolic strain at elevated pressure. Our data shows that PAH causes an increase in

both fresh tissue and purified elastin systolic stiffness by 510% and 123% (ΦSys, Cont to PAH)

respectively. By comparing the systolic and diastolic stiffness change from control to

hypertensive tissues (ΦSys, Cont to PAH to ΦDias, Cont to PAH), we see that the systolic stiffness

elevation of fresh tissue, resulting from PAH, is nearly double that of the diastolic stiffness

elevation, while the elastin stiffness elevation remains relatively constant. This is not surprising

given the linear nature of the elastin stress-strain curve compared to the highly non-linear fresh

17

tissue curve. This highly nonlinear behavior coupled with the elevated PAH physiologic strain

results in an increase in stiffness for control and PAH tissue from diastolic to systolic strain of

145% (ΦCont, Dias to Sys) and 278% (ΦPAH, Dias to Sys) respectively, while the more linear elastin

tissue only increased 39% and 14% for control and PAH tissues respectively. Also, the

minimum percent systolic load-carrying capacity of elastin is reduced from 48% (%Fn Sys, Cont)

for control tissues to 32% (%Fn Sys, PAH) for PAH tissues indicating that collagen carries more

load in the diseased state. However, these data also shows that elastin is an important load

carrying material even at PAH systolic strains. Increased systolic tissue stiffness is therefore due

to both elevated elastin stiffness as well as a higher degree of collagen engagement resulting

from the increase in systolic strain due to PAH.

While our results indicate that PAH causes PA tissues to operate at higher strains, the

elevated stiffness of the elastin dominant region mitigates this effect and maintains the

physiologic strain range within the transitional region. Due to the large strains associated with

control fresh tissues operating at PAH physiologic pressures we were unable to consistently

stretch the control fresh tissues to strains large enough to attain hypoxic pressure-load values

without risk of tissue damage. However, by projecting the fitted polynomial curves of control

fresh tissues, we estimate that the PAH diastolic pressure intersects the control fresh tissue curve

at a strain of ~73%. At these high strains, the compliance of arterial tissue is significantly

reduced due to elevated tissue modulus resulting from increased collagen loading. Therefore, the

elevated stiffness of the PAH elastin dominant region of the stress-strain curve maintains the

physiologic strains of PAH arterial tissues within the transition region of the stress-strain curve

by lowering the diastolic strain from ~73% to 53%. Without this increased stiffness, PAH PA

tissues would operate entirely within the high modulus, collagen dominant, region of the stress-

18

strain curve, which would result in a significant reduction of the Windkessel effect. This further

underlines the importance of understanding how elastin contributes to the mechanical properties

and morphology of PA tissue in response to PAH.

Pulmonary Vascular Stiffness and Elastin

Typically, modulus is calculated with a cross-sectional area determined from microscopic

analysis of arterial sections(8-10, 16, 27, 39, 42). This method assumes that the constituent

artery materials maintain homogeneity regardless of pathology or anatomical location. Since

stiffness calculations do not require this assumption, the 72% increase in elastin stiffness is an

independent result indicating that elastin carries substantially more load in the pathologic state.

However, the removal of non-elastin tissue during the digestion process results in experimentally

measured material properties that differ between the isolated elastin tissue and elastin in its

native state.

Individual elastic fibers are arranged in a network and bound together to form fenestrated

elastic sheets which surround the artery in concentric layers and work to uniformly distribute

arterial loads (38). While the individual elastic fibers do not have a linear stress-strain

relationship, network effects allow the resulting elastin tissue to obey Hooke’s law up to large

strains(45). The resulting elastic sheets, which are bound to non-elastin ECM components and

smooth muscle cells through glycoprotein and integrin bonds, imbue the artery with a Neo-

Hookean load-carrying elastin component that behaves in a near perfectly elastic manner(1).

Hypertensive adaptation of these intra-tissue bonding proteins is thought to play a role in the

development of increased arterial stiffness (2-4, 6, 20-22); however, during the CNBr digestion

process the elastin intra-tissue bonds are broken and all non-elastin material is extracted. In

addition, the removal of bulk, non-elastin, artery material eliminates many of the constraints

19

which resist the movement of elastin sheets relative to one another. The removal of both the

intra-tissue bonds and non-elastin material likely results in experimentally measured stiffness

and modulus values that are lower then those of the tissue in its native state. However, any

discrepancy between the experimentally-determined and native elastin stiffness should have little

impact on material property comparisons between control and hypertensive elastin tissues, since

both underwent the same digestion process. Due to this discrepancy between the experimentally

measured and native stiffness properties, %Fn-C and %Fn-H should be considered as the lower

bound for the fresh tissue load fraction carried by elastin and %Fn-e,f defines the minimum

fraction of increased load directly attributable to hypertensive elastin tissue. Hypertension,

therefore, causes a significant increase in the stiffness of the elastin material responsible for

carrying roughly half (%Fn-C, %Fn-H) of the arterial tissue load in the elastin dominant region, at a

minimum.

PVS is a function of both the intrinsic material property of modulus and the extrinsic

geometric consideration of thickness. Our data show a 46% and 28% average increase in fresh

tissue modulus and thickness respectively and an 81% average increase in fresh tissue stiffness.

For elastin tissues our data show a 65% and 6.9% average increase in modulus and ECF

respectively and a 72% average increase in stiffness. Since the width-normalized force, and

hence stiffness, scale linearly with changes in both modulus and thickness, it is evident that

modulus, rather than geometric tissue thickening, is the dominant factor in determining the

stiffness of hypertensive arterial and elastin tissues. Also, it is interesting to note that the

stiffness elevation predicted by our fresh tissue modulus and thickness data is 87%, which is

close to the measured stiffness increase of 81%, and the stiffness elevation predicted by our

20

increase in elastin modulus and ECF is 76%, which is close to the 72% average increase in

measured elastin stiffness.

Elastin Modulus

Our biomechanical tests show that hypertension causes an increase in the modulus of

both the fresh and elastin arterial tissues by an average of 46% and 65% respectively. This

increase in elastin modulus supports our thinking that one of the primary structural targets of

proximal vascular remodeling during hypoxia-induced PAH is arterial elastin. Changes in fresh

artery modulus are comparable to the 30% average increase in the modulus of the PAs in mice

published by Kobs et al (26, 27) and 40% decrease in compliance, identified as the threshold for

pediatric hypertension in humans, by Dyer et al (12).

The increased modulus of elastin, resulting from hypertension, suggests either a change

in elastin fibril modulus or a structural adaptation in fibril orientation / inter-elastin binding. The

turnover rate for collagen and elastin is low in healthy arteries, but vascular pathology upsets the

regulatory pathways that maintain this balance. In response to hypertension, the over expression

of both pro-inflammatory and proteinase-inhibitory molecules dramatically increases arterial

ECM synthesis (2, 23, 24). Although little is known about the ECM synthesis pathway of

hypertensive arterial tissue, (2, 40) the upregulation of the MMP-1 (matrix metalloproteinase -1)

inhibitor TIMP-1 (Tissue Inhibitor of Matrix Metalloproteinase 1) is one method by which ECM

synthesis may be enhanced during hypertension (23, 24). However, the ECM proteins

synthesized in response to hypertension have a three-dimensional architecture which is

functionally less optimal then those deposited during fetal development, and may play an

important role in determining the modulus of pathologic elastin tissue (2, 23, 24). More

21

experimentation is required to determine the underlying cause for the increase in elastin

modulus; this will be the topic of future study.

Limitations

Several limitations in these studies must be acknowledged. First, the tissue digestion

procedure may alter the mechanical properties of elastin through chemical attack or thermal

degradation, although no evidence of this was found in our studies or in the literature. Second,

the use of continuum mechanics for discontinuous materials will lead to some degree of error in

calculated mechanical properties due to simplifying assumptions; however, this should have little

impact on comparison of those properties. Third, uni-axial rather than fully biaxial testing was

performed. However, given that the objective of this study was not to investigate the anisotropic

behaviors of artery tissues, but rather to determine how such properties of fresh and digested

tissue samples change as a result of hypertension, the uni-axial test was deemed sufficient.

Conclusion

Via biomechanical testing of both fresh and digested elastin tissues of control and

hypoxia-induced pulmonary hypertensive animals, we have confirmed that vascular remodeling

results in a significant increase in the modulus and stiffness of both the fresh and elastin tissues

of the proximal PAs in the circumferential direction. We have further shown that, at 35% strain

a minimum 48% of the load imposed on fresh arterial tissue is carried by elastin, and that, at a

minimum, 43% of the change in stiffness of fresh arterial tissue is directly attributable to the

change in stiffness of the elastin component of those arteries. We have also shown that PAH

causes an increase in the strains associated with the physiologic pressure range but had no effect

on the strain of transition from elastin-dominant to collagen-dominant behavior. Stress-strain

22

behavior of fresh arterial tissues further revealed that the diastolic strain and baseline slope of the

stress-strain curve are determined by elastin mechanics, and the strain at systole becomes

increasingly dependent on collagen mechanics as a result of hypertension. The final conclusion

drawn from the above statements is that elastin is an important passive material for carrying the

load imposed on the proximal PAs at physiologic pressure ranges, and that the increase in

stiffness of elastin acts to maintain the physiologic strain range within the more elastin dependent

region of the stress-strain curve.

Acknowledgments

The authors would like to thank Dr. Conrad Stoldt for the use of laboratory equipment and

Colorado State University for providing access to their hypobaric chamber.

Sources of Funding

NIH P50HL84923

NIH T32HL072738

NIH K24HL081506

Disclosures

None

References

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26

Figure 1: Thickness measurements, tissue dissection and biopsy locations (typical). Internal Diameter (ID)

Figure 2: Detail of material testing system used to generate stress-strain data for arterial tissue samples.

Figure 3: Typical loading / unloading curves for fresh and elastin tissue tests.

Figure 4: Left: typical stress-strain behavior of fresh and elastin tissue, εtrans is the strain of transition from

the elastin-dominant region (A) to the transitional region (B) associated with increasing collagen engagement.

Right: typical curvature plot of fresh tissue.

Figure 5: Comparison of average thickness values and elastin area fractions between control and hypoxic

populations. (* P<0.05)

Figure 6: Comparison of mean values for artery stiffness and modulus of the RPA, LPA and MPA for control

and hypoxic populations. (* P<0.05)

Figure 7: Average control and hypertensive tissue stiffness at systolic and diastolic pressures. Comparative

values presented in table 1.

Figure 8: Comparison of mean values for elastin stiffness and modulus for the RPA, LPA and MPA for

control and hypoxic populations. (* P < 0.05)

Figure 9: Mean pressure-strain data for control and hypertensive tissues. Shaded regions indicate the

predicted physiologic conditions for control (A) and hypertensive (B) tissues.

Table 1: Average change in stiffness for control and PAH tissues at diastolic and systolic pressure and

average load carrying capacity of elastin. Comparative values based on the data presented in figure 7. ΦDias

27

(Sys), Cont to PAH: average increase in stiffness from control to PAH tissues measured at diastole (systole). ΦCont

(PAH), Dias to Sys: average increase in control (PAH) tissue stiffness from diastole to systole. %Fn Dias (Sys), Cont (PAH):

average percent load carrying capacity of control (PAH) tissue measured at diastole (systole). († P < 0.05)

ID-1

ID-2

Circumferential

MC3

ML3

MC2

ML2

MC1M

L1

Spare

Tissue anomaly (avoided)

Sec

tioni

ng P

lane

Thickness Measurement 4-Ortho. ea. side (8 total)

Axial View

Transverse Iso. View Sectioning Diagram

Biopsy tissueMech. test samples

5N Load Cell

A

A

Section AAMTS Detail

Buffer Solution

Rec

ycle

Loo

p

Fluid Level

5N L

oad

Cel

l

Thumb Screw

TissueGrips

ArterialTissueSample

Environmental Chamber Wall

−20 0 20 40 60 80−100

0

100

200

300

400

Strain (%)

Load

(m

N)

Fresh Tissue Raw Data, Fn − Strain Typ.

Preconditioning

Conditioned

−50 0 50 100 150−20

0

20

40

60

80

100

120

Strain (%)

Load

(m

N)

Elastin Tissue Raw Data, Fn − Strain Typ.

0 20 40 600

50

100

150

200

Strain ε (%)

Str

ess

σ (k

Pa)

Stress vs. Strain (Typ.)

Diastolic Stress

Systolic Stress

A B

σfresh

σelast

0 20 40 60 800

1

2

3

4

5x 10

−9 Strain vs. Curvature (Typ.)

Strain ε (%)

Cur

vatu

reκ

κAverage κ

PSS

εtrans

εL

εH

εPSS

H

κmax

κtrans

RPA LPA MPA0

1

2

3

4

5

6Average Thickness

Tissue Type

Thi

ckne

ss(m

m)

* *

*Control

Hypoxic

RPA LPA MPA0

10

20

30

40

50Average Area Fraction

Tissue Type

Are

a F

ract

ion

(%) * * *

RPA LPA MPA0

100

200

300

* * *

Artery Stiffness: Circumferential Direction

Tissue Type

Stif

fnes

sΦ

(m

m k

Pa)

RPA LPA MPA0

50

100

150

200

*

*

Artery Modulus: Circumferential Direction

Tissue Type

Mod

ulus

E (

kPa)

Control

Hypoxic

RPA LPA MPA0

1000

2000

3000

4000Average Artery Stiffness for Systolic and Diastolic Pressures

Tissue Type

Stif

fnes

sΦ

(m

m k

Pa)

RPA LPA MPA0

100

200

300

400Average Elastin Stiffness for Systolic and Diastolic Pressures

Tissue Type

Stif

fnes

sΦ

(m

m k

Pa)

Diastolic Control Diastolic Hypoxic Systolic Control Systolic Hypoxic

RPA LPA MPA0

50

100

150

200

* *

*

Elastin Stiffness: Circumferential Direction

Tissue Type

Stif

fnes

sΦ

(m

m k

Pa)

RPA LPA MPA0

50

100

150

200

* *

Elastin Modulus: Circumferential Direction

Tissue Type

Mod

ulus

E (

kPa)

Control

Hypoxic

RPA LPA MPA Average RPA LPA MPA Average

Dias, Cont to PAH 327%† 274%† 279%† 293%† 151%† 243%† 124%† 173%†

Sys, Cont to PAH 584%† 512%† 435%† 510%† 102%†186% 81%† 123%†

Cont, Dias to Sys 151%† 125%†158% 145%†

41% 41% 36% 39%

PAH, Dias to Sys 303%† 268%† 265%† 278%†13% 18% 10% 14%

RPA LPA MPA Average%Fn Dias, Cont 50% 46% 87% 61%%Fn Dias, PAH 45% 43% 56% 48%%Fn Sys, Cont 45% 36% 63% 48%%Fn Sys, PAH 30% 31% 35% 32%

Table 1. Physiologic stiffness comparison Fr

esh

Tiss

ue

Ela

stin

Tis

sue

Percent load carrying capacity of elastin