Maladaptive aortic remodeling in hypertension associates with dysfunctional smooth muscle contractility

Abstract

Intramural cells are responsible for establishing, maintaining, and restoring the functional capability and structural integrity of the aortic wall. In response to hypertensive loading, these cells tend to increase wall content via extracellular matrix turnover in an attempt to return wall stress and/or material stiffness toward homeostatic values despite the elevated pressure. Using a common rodent model of induced hypertension, we found marked mouse-to-mouse differences in thoracic aortic remodeling over 2–4 wk of pressure elevation, with mechanoadaptation in some but gross maladaptation in most mice despite the same experimental conditions and overall genetic background. Consistent with our hypothesis, we also found a strong correlation between maladaptive aortic remodeling and a dysfunctional ability of the vessel to vasoconstrict, with maladaptation often evidenced by marked adventitial fibrosis. Remarkably, mouse-to-mouse variability did not correlate with the degree or duration of pressure elevation over the 2- to 4-wk study period. These findings suggest both a need to study together the structure, mechanical properties, and function across layers of the wall when assessing aortic health and a need for caution in using common statistical comparisons across small seemingly well-defined groups that may mask important underlying individual responses, an area of investigation that demands increasing attention as we move toward an era of precision diagnosis and patient care.

NEW & NOTEWORTHY There are three primary findings. Marked mouse-to-mouse differences exist in large vessel hypertensive remodeling in an otherwise equivalent cohort of animals. The degree of maladaptation correlates strongly with decreases in smooth muscle contractile capacity. Finally, short-term maladaptive remodeling is independent of the precise degree or duration of the pressure elevation provided that thresholds are exceeded. Therapeutic targets should thus be personalized and focus on both layer-to-layer interactions and early interventions.

INTRODUCTION

Systemic hypertension is thought both to cause and to be caused by central artery stiffening due to complex feedback loops between the local mechanics/mechanobiology and global hemodynamics/physiology (25, 28). Identifying biomechanical factors underlying such stiffening as well as approaches for preventing or reversing the associated maladaptive changes in aortic geometry, wall structure, and overall function is thus a critical area of investigation (6, 33). A vast literature reveals, for example, marked changes in extracellular matrix composition and passive mechanical properties of the aorta in uncontrolled hypertension (5, 21, 22). Consistent with these reports, we recently observed a dramatic maladaptation of the thoracic aorta in a mouse model of induced hypertension that is characterized largely by an inflammation-driven fibrosis of the adventitia (4). There has been much less investigation of the effects of or potential changes in smooth muscle contractility, however. Hence, we used this same mouse model to test the hypothesis that aortic maladaptations in hypertension associate with a compromised vasoconstriction capability.

Specifically, we quantified and contrasted contributions of extracellular matrix remodeling and smooth muscle dysfunction to the maladaptive morphology and biomechanical responses of the proximal descending thoracic murine aorta to angiotensin II (ANG II)-induced hypertension. To examine possible differences due to the duration of short-term hypertension, we compared the degree of changes in matrix remodeling and smooth muscle contractility over 2–4 wk of ANG II infusion. Although neither the duration of infusion nor precise level of pressure elevation correlated with remodeling over this period, marked differences manifested among the animals. Despite preserved elastic lamellae and only modest changes in the overall intrinsic passive circumferential stiffness of the wall, cases of moderate-to-severe maladaptation associated with a marked deposition of medial and especially adventitial collagen, the latter consistent with a layer-specific fibrotic thickening. This increase in collagen content was likely responsible for a decreased elastic energy storage capability of the wall, that is, a significantly reduced mechanical functionality. Despite only modest decreases in smooth muscle α-actin and myosin heavy chain density, the cases of maladaptive remodeling also associated with marked reductions in vessel-level vasoconstriction capacity, consistent with compromised medial smooth muscle contractility. Importantly, comparing changes in passive and active functionality revealed a strong correlation between maladaptive wall remodeling and diminished vasoconstriction capacity. These findings suggest an, at least initially preserved, ability of smooth muscle cells to mechanosense and mechanoregulate matrix within the media via actomyosin activity despite a loss of vessel-level vasoconstriction function and emphasize the importance of considering together any changes in matrix and smooth muscle phenotype.

MATERIALS AND METHODS

Mouse model.

All animal procedures were approved by the Institutional Animal Care and Use Committee of Yale University. To minimize extrinsic variability, we used adult wild-type mice (C57BL/6J, Jackson Laboratories) of the same sex (male) maintained on the same diet (normal) and exposed to the same light-dark cycles and limited exercise. Specifically, hypertension was induced in 12- to 24-wk old mice via chronic infusion of ANG II (catalog no. A9525, Sigma-Aldrich) at a continuous rate of 490 ng·kg⁻¹·min⁻¹ delivered via a subcutaneously implanted osmotic minipump (model 2004, Alzet). Conscious resting blood pressures were measured noninvasively using a CODA tail-cuff system (Kent Scientific, Torrington, CT) following standard methods. At the predetermined end point, mice were euthanized via an intraperitoneal injection of Beuthanasia-D (150 mg/kg body mass, catalog no. 1848, Henry Schein) and exsanguination upon transection of the descending thoracic aorta (DTA). The duration of ANG II infusion ranged from 2 to 4 wk to assess possible effects of infusion time: n = 2 mice for each day between 13 and 22 days with an additional mouse (n = 1) at 15, 27, and 28 days. Noninfused mice, littermates where possible, served as controls. Additional details on this induced model of hypertension can be found elsewhere (4, 43).

Specimen preparation.

Within 15 min of death, the proximal DTA, from the left subclavian artery to the third pair of intercostal branches, was excised en bloc and placed in Krebs-Ringer solution at room temperature. Specimens were cleaned of excess perivascular tissue by gentle dissection under a dissection microscope, cannulated on custom-drawn glass micropipettes (P-97 Flaming/Brown Micropipette Puller, Sutter Instruments), secured with 6-0 suture at each end, and then mounted within a custom computer-controlled biaxial testing system (4, 14). After being tested, specimens were unloaded, fixed overnight in 10% neutral buffered formalin, and then stored in 70% ethanol at 4°C until all samples could be processed together for histological examination.

Active wall mechanics.

The general procedure for assessing active biaxial wall mechanics has been described elsewhere (15, 30). Briefly, the testing chamber was filled with phosphate-buffered Krebs-Ringer solution that was heated slowly to and then maintained at 37°C while bubbled with 95% O₂-5% CO₂ to maintain pH at 7.4. After the unloaded configuration was recorded, vessel viability was assessed as responsiveness of smooth muscle contraction by increasing KCl to 80 mM at a modest luminal pressure (P = 40 mmHg) and axial stretch (λ_z = 1.1) followed by smooth muscle relaxation in normal Krebs solution and then a second contraction at P = 60 mmHg and λ_z = 1.2. After the second relaxation, the in vivo (or preferred) passive value of axial stretch, ${\text{λ}}_z^{\text{iv}}$, was estimated as previously described (4), namely, by axially extending the vessel in steps until cycles of pressurization from ~80 to 100 mmHg associated with an almost constant axial force measurement. After the vessel was set at this axial stretch, smooth muscle contractions to high K⁺ were elicited at seven to nine combinations of transmural pressure (e.g., 70, 90, or 110 mmHg) and axial stretch (${\text{λ}}_z>{\text{λ}}_z^{\text{iv}},{\text{λ}}_z\le {\text{λ}}_z^{\text{iv}},{\text{λ}}_z\ll {\text{λ}}_z^{\text{iv}}$) for 10–15 min each, with 10–30 min periods of relaxation (i.e., washout of high K⁺) between each contraction experiment. After each change in pressure or axial stretch, the specimen was allowed to equilibrate under nonstimulated conditions (regular Krebs-Ringer solution) for 5–10 min before the next stimulation with high K⁺. This range of pressure-stretch combinations was identified via pilot studies focusing on maintaining smooth muscle viability during biaxial loading. Two vessels from ANG II-infused mice responded poorly to initial smooth muscle contraction and relaxation. The active wall mechanics protocol was suspended after the data were recorded from one or two pressure-stretch combinations for these two specimens.

Percent changes in outer diameter, axial force, and associated biaxial stresses were determined as 100 × (value prestimulus − value poststimulus)/(value prestimulus). Mean values of circumferential and axial stresses (σ_θ and σ_z, respectively), each of which depend on the mean values of circumferential and axial stretches (λ_θ and λ_z, respectively), were calculated using the following standard formulas (22),

$${\text{σ}}_{\text{θ}}\left({\text{λ}}_{\text{θ}},{\text{λ}}_z\right)=\frac{\text{P}a}{h}$$ (1)

$${\text{σ}}_z\left({\text{λ}}_{\text{θ}},{\text{λ}}_z\right)=\frac{f_{\text{T}}+\text{π}{a}^2\text{P}}{\text{π}h\left(2a+h\right)}$$ (2)

where P is the transmural pressure measured with standard transducers, a is the inner deformed radius, h is the deformed thickness, and f_T is the axial load measured directly by a force transducer. Both a and h were calculated by assuming incompressibility during loading and contraction using the unloaded volume coupled with online measurements of outer diameter and axial length, which allowed the calculation of biaxial stretches using standard formulas (i.e., ratios of current to unloaded mean radius and length, respectively).

Passive wall mechanics.

After the active testing was completed, the vessel was rendered noncontractile via immersion in and pressurization with HBSS (catalog no. 14025-134, GIBCO) as previously described (14). All passive biaxial mechanical testing and data analysis were then performed consistent with previousy described methods (4, 14). Briefly, the passive unloaded configuration was determined with the outer diameter measured using a video microscope and wall thickness measured using an optical coherence tomography system (ThorLabs, Newton, NJ). Next, vessels were mechanically preconditioned by subjecting them to cyclic pressurization while held at their preferred (in vivo) axial stretch (${\text{λ}}_z^{\text{iv}}$) and then subjected to a series of seven biaxial protocols: cyclic pressurization from ~10 to 140 mmHg at ${\text{λ}}_z^{\text{iv}}$ and repeated at ±5% of this stretch as well as cyclic axial stretching at constant luminal pressures of 10, 60, 100, and 140 mmHg.

The associated passive pressure-diameter and axial force-length data were then fit with a validated four-fiber family constitutive model (2, 4, 15) via a nonlinear regression of a data set consisting of the final cycle of unloading data from all seven testing protocols. The data during unloading revealed the elastic energy stored during deformation that would be available to work on the distending fluid. This constitutive relation includes eight model parameters within a Holzapfel-type nonlinear stored energy function (W), namely:

$$W\left(C,M^i\right)=\frac{c}{2}\left(I_C-3\right)+{\displaystyle \sum _{i=1}^4\frac{c_1^i}{{4c}_2^i}\left\{\exp \left[c_2^i{\left(I{V}_C^i-1\right)}^2\right]-1\right\}}$$ (3)

where c, $c_1^i$, and $c_2^i$ (i = 1, 2, 3, or 4 for the four predominant fiber family directions) are material parameters, with c and $c_1^i$ having units of stress (in kPa) and $c_2^i$ as dimensionless. I_C = tr(C) and $I{V}_C^i=M^i\hspace{0.17em}·\hspace{0.17em}C{M}^i$ are coordinate invariant measures of the finite deformation, with the right Cauchy-Green tensor C = F^TF computed from the deformation gradient tensor F = diag[λ_r, λ_θ,λ_z], with detF = 1 because of assumed incompressibility. The direction of the ith family of fibers was identified by the vector $M^i=\left[0,\sin {\alpha }_0^i,\cos {\alpha }_0^i\right]$, with the model parameter ${\alpha }_0^i$ denoting a fiber angle relative to the axial direction in the traction-free reference configuration. Based on prior microstructural observations from multiphoton microscopy and second harmonic generation (14), and because of the yet unknown effects of copious cross links among the multiple families of fibers, the four predominant fiber families were axial (${\alpha }_0^1=0$), circumferential (${\alpha }_0^2=\text{π}/2$), and symmetric diagonal (${\alpha }_0^{3,4}=\pm {\alpha }_0$). Values of mean biaxial wall stress and material stiffness were computed from the stored energy function and calculated at individually measured values of systolic or mean blood pressure as well as at a common pressure of 100 mmHg.

Histology and immunohistochemistry.

Fixed samples were dehydrated, embedded in paraffin, sectioned serially (5 μm thickness), and stained with Verhoeff Van Gieson (VVG), Masson’s trichrome (MTC), or picrosirius red (PSR) for standard histology or an antibody for CD45 (Mouse LCA), a pan inflammatory cell marker. Detailed analyses were performed on two biomechanically representative vessels for each of four primary “groups” (see below). Custom MATLAB scripts (4) were used to extract layer-specific cross-sectional areas and to calculate the positively stained CD45⁺ pixels and later immunofluorescent-positive pixels. Colorimetric analyses of PSR-stained cross sections revealed four subtypes of fibrillar collagen in the adventitia, indicated by red, orange, yellow, and green birefringent fibers under polarized light. Medial-adventitial percentages were averaged from MTC-stained images and confirmed with VVG-stained images. Histological images were acquired on an Olympus BX/51 microscope (under normal or polarized imaging) using an Olympus DP70 digital camera (CellSens Dimension) and a ×20 magnification objective.

For immunofluorescence staining, cut sections were deparaffinized and rehydrated before heat-mediated antigen retrieval in citrate buffer for 25 min. After being blocked with Background Sniper (Biocare Medical) to reduce nonspecific background staining, sections were incubated overnight with antibodies for smooth muscle α-actin (ab5694, Abcam, 1:100) or smooth muscle myosin heavy chain (ab53219, Abcam, 1:100). Primary antibody binding was visualized using species-specific Alexa Fluor 594-conjugated secondary antibodies (Invitrogen). Sections were mounted using ProLong Gold antifade reagent with DAPI (Invitrogen). Each tissue section was imaged under three separate fluorescent channels: TRITC for the molecular marker of interest, green to capture elastin autofluorescence, and DAPI for cell nuclei. The media and adventitia were delineated based on the location of the external elastic lamellae in the green elastin autofluorescent images. Immunofluorescent images were acquired on a Nikon Eclipse Ti-S microscope using a ×20 magnification objective.

Statistics.

We fit data on maximal vessel-level constrictions [i.e., maximum change in outer diameter, both absolute (in μm) and as a percentage, with n = 5 noninfused control mice and n = 23 ANG II-infused mice] to a normal distribution function using the normfit function (MATLAB, MathWorks) and reported the mean (μ) and standard deviation (σ). Tukey’s rule (the interquartile range) did not identify any data outside the expected range; hence, all data were included in the final analysis. We then delineated observations that were 1) well above the mean for the infused group (i.e., above μ + σ), which we refer to these active responses as near normal, similar to those of control aortas; 2) close to the mean of the infused group (within μ ± σ), which we refer to as a compromised active response to high K⁺; and 3) well below the mean (i.e., below μ – σ), which we refer to as an impaired active response. To determine possible correlations between the ability of the vessel to vasoconstrict and the passive mechanical metrics, we computed Pearson’s linear correlation coefficient and the corresponding P value for ANG II-infused specimens. Correlations were calculated for a data matrix containing a vasoconstriction parameter (maximum percent change in outer diameter) as well as passive geometric, material, and structural metrics (using the corr function, MATLAB, MathWorks). All data are presented as means ± SE unless otherwise noted. Differences between the noninfused control group and the overall ANG II-infused group were assessed using a t-test with Welch’s correction for unequal variances (GraphPad Prism 7). P < 0.05 was considered significant.

RESULTS

ANG II induces hypertension and overall maladaptive remodeling.

Infusion of ANG II (490 ng·kg⁻¹·min⁻¹) raised resting tail-cuff measured systolic blood pressure by 45 mmHg (from 125 ± 8 to 170 ± 14 mmHg, that is, a 36% increase) and diastolic blood pressure by 30 mmHg (from 92 ± 19 to 122 ± 13 mmHg, a 33% increase), both increases being statistically significant (Welch’s t-test, P < 0.0001). Importantly, the specific elevation in blood pressure varied from mouse to mouse but was otherwise individually maintained at essentially the same level from 2 to 4 wk, the period of study. This degree of pressure elevation and its near steady state over 2–4 wk is consistent with previously reported ambulatory telemetric values, as, for example, systolic pressure P_sys = 121 mmHg at baseline and P_sys = 167 mmHg after 2 wk of ANG II infusion at the same dose (43). Also consistent with prior observations (4), hypertension led to an overall maladaptive geometric remodeling of the DTA. On average, total wall thickness at systolic pressure increased by ~1.97-fold. In contrast, a mechanoadaptive remodeling with a constant cardiac output should have increased wall thickness only 167/121 = 1.38-fold (24).

Standard passive biaxial data (n = 5 controls and n = 23 ANG II infused) suggested that ANG II infusion alters the overall structural behavior primarily in the axial direction (Fig. 1, A and B). Quantification of these data using the four-fiber family stored energy function (Table 1) revealed further that associated biaxial mechanical, not just geometric, changes were independent (P > 0.05) of both the duration of ANG II infusion (from 2–4 wk) and the precise degree of pressure elevation (P_sys = 170 ± 14 mmHg) across the 23 hypertensive mice (Fig. 2). Albeit not shown, passive results were also independent of the initial age of the mice over the narrow range studied (12–24 wk). Figure 1, C and D, also shows mean circumferential stress-stretch and axial stress-stretch behaviors under passive conditions. These data suggested that there was little change in material stiffness despite the overall structural stiffening (leftward shifts in Fig. 1, A and B), although with a reduction in biaxial wall stress (P < 0.01) that likely resulted from the combined effects of an excessive increase in wall thickness (from 38 to 75 μm at P_sys, P < 0.01) and a marked decrease in the in vivo axial stretch (from 1.53 to 1.45, P < 0.05). Detailed calculations at group-specific P_sys values are shown in Table 2, where it can also be seen that the ability of the hypertensive vessels to store elastic energy decreased significantly. Loss of energy storage capability reveals a diminished mechanical functionality since an important function of elastic arteries is to store energy during systole that can be used during diastole to work on the blood and augment flow. Figure 3 shows the various metrics for all samples at a common pressure of 100 mmHg to avoid any pressure effect. A comparison of the results between Table 2 and Fig. 3 revealed two important observations: first, that biaxial material stiffness appeared to be adapted to the elevated pressure since these values are lower than normal, on average, in ANG II-infused samples when compared at 100 mmHg and, second, that there was considerable variability in the ANG II-infused group, with some specimens exhibiting near normal wall thickness, axial stretch, energy storage, and wall stress when compared at the same pressure.

Fig. 1.

Mean passive mechanical testing data for the proximal descending thoracic aorta (DTA) from male wild-type mice. A–D: data are from noninfused (control, black circles, n = 5) and ANG II-infused (duration of 13–28 days, white circles, n = 23) mice. E–H: shown, too, are the same mean noninfused (black circles, n = 5) data compared against data for 3 ANG II-infused samples (n = 1 each), which are representative of vessels that had a near normal (norm, dark gray circles), compromised (comp, light gray squares), or impaired (imp, white inverted triangles) ability to vasoconstrict in response to 80 mM KCl stimulation. Structural responses were revealed by pressure-diameter data at group-specific in vivo axial stretches (A and E) and axial force-axial stretch data (B and F) at 100 mmHg. Material responses were revealed by mean circumferential (Circ.) Cauchy stress-stretch data at in vivo axial stretches (C and G) and axial Cauchy stress-stretch data (D and H) at 100 mmHg. Bars (means ± SE) are barely evident in most of curves in A–D because of either low variability (controls) or a large sample size (infused).

Table 1.

Best-fit values of model parameters in the strain energy function W (Eq. 3) for the DTA, determined from the seven different passive pressure-diameter and axial force-length protocols, based on individual results for all specimens

	Elastic Fibers	Axial Collagen		Circumferential Collagen + Smooth Muscle Cells		Symmetric Diagonal Collagen			Error
	c, kPa	$c_1^1$, kPa	$c_2^1$	$c_1^2$, kPa	$c_2^2$	$c_1^{3,4}$, kPa	$c_2^{3,4}$	αο,°	Root Mean Square Error
Control (n = 5)	26.73 ± 3.10	20.82 ± 2.04	0.25 ± 0.08	12.74 ± 0.94	0.10 ± 0.03	0.22 ± 0.12	2.19 ± 0.40	28 ± 2	0.086
ANG II infused
All (n = 23)	16.50 ± 2.11	11.54 ± 2.11	1.32 ± 0.30	9.21 ± 0.91	0.88 ± 0.42	3.98 ± 2.57	4.28 ± 0.92	32 ± 2	0.121
Normal (n = 3)	27.84 ± 7.17	11.09 ± 2.31	0.34 ± 0.15	8.30 ± 0.88	0.13 ± 0.02	0.37 ± 0.21	2.00 ± 0.83	27 ± 2	0.065
Compromised (n = 16)	15.08 ± 2.38	12.21 ± 2.86	1.35 ± 0.37	8.94 ± 0.94	0.55 ± 0.20	5.60 ± 3.65	3.21 ± 0.75	32 ± 2	0.111
Impaired (n = 4)	13.69 ± 3.26	9.18 ± 4.39	1.90 ± 0.86	10.98 ± 3.97	2.73 ± 2.28	0.19 ± 0.18	10.23 ± 3.05	34 ± 3	0.202

Values are means ± SE. Specimens were from the noninfused control group (n = 5) and ANG II-infused group (n = 23), with the latter also subdivided based on contractile capacity (near normal, compromised, or impaired; see Fig. 4). Because the constitutive model is microstructually motivated but phenomenological, one should be cautious when interpreting changes in individual parameters across groups; it is better to focus on integrative effects calculated in terms of wall stress, stiffness, and energy storage (see Fig. 6 and Table 2).

Fig. 2.

Values of passive geometric and mechanical metrics for the descending thoracic aorta (DTA) from all n = 23 ANG II-infused mice computed at a common pressure of 100 mmHg but individual values of in vivo axial stretch and then plotted as a function of the duration of ANG II infusion (A–D) or mouse-specific systolic pressure (E–H). Data were further distinguished based on their ability to vasoconstrict in response to 80 mM KCl (cf. Fig. 4) as follows: near normal (dark gray circles, n = 3), compromised (light gray squares, n = 16), and impaired (white inverted triangles, n = 4). Neither the duration of infusion nor systolic pressure correlated with DTA deformed thickness, circumferential (Circ.) stress or stiffness, or stored energy; although not shown, the same lack of correlation held for deformed outer diameter, axial stretch, and axial stress and stiffness (i.e., P > 0.05 and Pearson’s linear correlation coefficient |ρ| < 0.5).

Table 2.

Passive geometric, material, and structural metrics for noninfused, control and ANG II-infused specimens

	Not Infused (n = 5)	ANG II Infused (n = 23)
Unloaded dimensions
Outer diameter, µm	932 ± 21	1,059 ± 25*
Wall thickness, µm	108 ± 3.4	174 ± 6.9†
Inner radius, µm	359 ± 11.8	355 ± 11.5
In vitro axial length, mm	4.99 ± 0.31	4.92 ± 0.18
Systolic dimensions	P = 121 mmHg	P = 167 mmHg
Outer diameter, µm	1,556 ± 37	1,579 ± 31
Wall thickness, µm	38 ± 0.6	75 ± 6.0†
Inner radius, µm	740 ± 18.3	715 ± 20.0
In vivo circumferential stretch, λϐ	1.84 ± 0.03	1.72 ± 0.06
In vivo axial stretch, ${\text{λ}}_z^{\text{iv}}$	1.53 ± 0.04	1.45 ± 0.03
Systolic Cauchy stresses, kPa
Circumferential, σθ	312 ± 6.0	255 ± 25.6*
Axial, σz	266 ± 13.3	187 ± 21.2*
Systolic linearized stiffness, MPa
Circumferential, Cθθθθ	2.23 ± 0.15	2.58 ± 0.16
Axial, Czzzz	3.99 ± 0.28	3.34 ± 0.47
Systolic stored energy, kPa	78 ± 3.9	50 ± 6.6*

Values are means ± SE. Pressure-dependent values were calculated at ambulatory normotensive (for control) or hypertensive (ANG II-infused specimens) systolic arterial pressures, where specified. Statistical significance was assessed by a t-test with Welch's correction for unequal variances. Because of subgroup sizes, statistical comparisons were not performed for near normal, compromised, and impaired ANG II-infused specimens. See Fig. 6 instead. The geometric values, in combination with the constitutive relation and best-fit parameters of Table 1, allow one to recreate the biaxial data if desired.

^*P < 0.05;

^†P < 0.001.

Fig. 3.

Comparison of geometric and material metrics between noninfused (left, n = 5) and ANG II-infused (right, n = 23) specimens. Values were calculated at a common pressure of 100 mmHg but individual values of in vivo axial stretch. Statistical significance was assessed using a t-test with Welch’s correction for unequal variances. Circ., circumferential. *P < 0.05, statistical significance.

Smooth muscle contractility is impaired in some, but not all, ANG II-infused mice.

Induced hypertension associated with an overall reduced capacity of the DTA to vasoconstrict in response to 80 mM KCl (Fig. 4A): the maximum reduction in outer diameter at 15 min in response to high K⁺ isobaric/axial isometric testing was 23.5 ± 1.0% in control specimens but only 16.41 ± 6.5% in ANG II-infused specimens (P < 0.05), a 30% decrease. It should be noted that each vessel had a unique combination of pressure-axial stretch that yielded its maximum constriction; results were thus compared for mouse-specific values of pressure and stretch to focus on maximum contractile capacity. Importantly, similar to findings for the passive properties, differences in vasoconstriction were independent (P > 0.05) of the initial age of the mice, the duration of ANG II infusion over the 2- to 4-wk period of interest, and the mouse-specific value of P_sys. Notwithstanding the observed difference in mean behaviors, the range of maximum change in diameter was narrow for control vessels (21–26%, with a standard deviation of 2%) but broad for ANG II-infused vessels (3–28%, with a standard deviation of 6.5%). In other words, there were ANG II-infused DTAs whose ability to vasoconstrict was impaired dramatically (n = 4), somewhat (n = 16), or not at all (n = 3) relative to that of the controls (n = 5), an observation that was independent of the specimen-specific value of pressure-axial stretch at which the vasoconstriction was maximal. These differences were revealed well by a frequency plot (Fig. 4B), which suggested a delineation of ANG II-infused specimens as follows: specimens with impaired (<μ – σ = 124 μm or <10% change in diameter), compromised, or near normal (>μ + σ = 305 μm or >24% change in diameter) smooth muscle contractility. As shown, 17% (4 of 23 specimens) presented as impaired, 70% (16 of 23 specimens) presented as compromised, and 13% (3 of 23 specimens) presented with near normal contractility in response to high K⁺ stimulation. Interestingly, the same pressure but mouse-specific axial stretches distended and extended, respectively, the different vessels to similar outer diameters. Figure 4C shows illustrative extremes in the observed vasoconstriction responses (i.e., diameter reduction and axial force as a function of time during isobaric-isometric testing), namely, one representative specimen from the control group and one representative specimen with impaired contractility.

Fig. 4.

A and B: maximum change (Δ) in outer diameter (OD) of the descending thoracic aorta (DTA) upon stimulation with 80 mM KCl at specimen-specific optimal combinations of pressure and axial stretch. All ANG II-infused (n = 23) specimens were compared with noninfused control specimens (n = 5). A: the percent change in OD was calculated as 100 × (OD prestimulus − OD poststimulus)/(OD prestimulus). Statistical significance was assessed by a t-test with Welch’s correction for unequal variances (*P < 0.001). B: histogram of the maximum change in diameter calculated as follows: (OD prestimulus – OD poststimulus). A normal distribution fit well the observed variations exhibited by the ANG II-infused group (dashed line); the associated fit for specimens in the control group (blue shaded curve) are plotted for comparison. The mean (μ) of both groups is annotated, and the means ± standard deviation (μ ± σ) of the ANG II-infused group are annotated to delineate specimens with impaired (white, < μ − σ), compromised (light gray, μ ± σ), or near normal (dark gray, > μ + σ) abilities to vasoconstrict. C and D: representative tracings of the online measurement of OD and axial force upon stimulation with 80 mM KCl shortly after 0 min for one ANG II-infused specimen with impaired contractility (in gray); shown for comparison are representative data from the control group (in black). Note the biaxial consequences of KCl-stimulated smooth muscle contractility in the control vessels.

Reduced contractile capacity associates with increased wall remodeling in hypertension.

Given the important finding of highly variable contractile responses across the 23 infused mice after duration-independent ANG II infusion for 2–4 wk, we reassessed the associated passive geometric and mechanical metrics in terms of the following same four “groups”: noninfusion control versus ANG II infusion having near normal, compromised, or impaired vasoconstriction capacity. Figure 5 shows representative histological cross sections for these groups. VVG-stained sections (Fig. 5, row 1) revealed intact elastic laminae in all groups but increasingly greater intralaminar distances in hypertensive specimens having decreased contractile capability. MTC-stained sections (Fig. 5, row 2) further suggested decreases in smooth muscle and increases in collagen in the media of hypertensive specimens, consistent with the increased intralaminar distances. In particular, the medial area increased 1.5-fold in the ANG II-infused group relative to the control group (Fig. 5B). These MTC-stained sections also showed that the increased adventitial thickness resulted largely from increased collagen deposition, especially in samples with an impaired ability to vasocontrict (Fig. 5A). Adventitial area increased by 1.3-fold in the representative samples having a near normal ability to vasoconstrict but over 2-fold and over 3-fold in the compromised and impaired vasoconstriction samples, respectively. The PSR-stained sections (Fig. 5, row 3) revealed further that the additional adventitial collagen contained an increased percentage of thinner fibers: 6.4- and 7.2-fold increases in yellow and green fiber area fractions, respectively, for the impaired vasoconstriction group relative to the control group (Fig. 5F). Note, too, the presence of smooth muscle α-actin-positive cells in the adventitia in the impaired vasoconstriction group, which were likely myofibroblasts. Overall, therefore, the impaired contractility group showed the most dramatic matrix remodeling, both medial and adventitial, whereas the near normal contractility group, although thickened, otherwise appeared similar to the control group. Immunofluorescent imaging of smooth muscle α-actin and smooth muscle myosin heavy chain (Fig. 5, rows 4 and 5) suggested that there was a progressive, although modest, decrease in smooth muscle contractile proteins with increasingly maladaptive remodeling.

Fig. 5.

A: computed area fractions of cross sections of the aortic wall that were medial (white) or adventitial (gray). B–E: shown, too, are cross sections for one representative vessel per subgroup [top to bottom: Verhoeff Van Gieson stain (VVG), Masson’s trichrome stain (MTC), and picrosirius red stain (PSR)] and merged immunofluorescent stains against smooth muscle myosin heavy chain or smooth muscle α-actin in red, with DAPI in blue. The edge of the adventitia is outlined with a white dashed line in rows 4 and 5. Values were compared between noninfused (control; B) and ANG II-infused (C–E) specimens grouped based on their ability to vasoconstrict in response to 80 mM KCl (recall Fig. 4) as follows: near normal (C), compromised (D), and impaired (E). F: list of computed cross-sectional area fractions in the adventitia for the red-, orange-, yellow-, and green-stained collagen fibers from polarized light imaging of PSR-stained sections (mean of 2 representative samples/group). Overall, note the slight decrease in medial smooth muscle myosin heavy chain and α-actin with decreased contractile capacity as well as the marked increase in adventitial collagen and myofibroblasts (white arrows, bottom right) in the impaired contractility specimen. Scale bar = 100 μm.

Given the broad range of effects of ANG II infusion on vasoconstriction, Fig. 1, E–H, delineates illustrative passive behaviors that correspond to representative specimens from the control group as well as the near normal, compromised, and impaired vasoconstriction groups. Importantly, Fig. 6 shows plots of mouse-specific passive geometric and mechanical metrics (calculated at 100 mmHg) as a function of the associated vasoconstriction capacity (i.e., maximum change in outer diameter upon high K⁺ stimulation). Plotted in this way, strong correlations emerged between the ability to vasoconstrict and passive outer diameter (Fig. 6A), wall thickness (Fig. 6E), in vivo axial stretch (Fig. 6B), stored energy (Fig. 6F), and circumferential (Fig. 6C) and axial (Fig. 6G) wall stress. In contrast, axial material stiffness correlated only mildly and circumferential material stiffness did not correlate well with contractile capacity, and these findings were even more definitive when stiffness was calculated at group-specific P_sys values (Table 2). It has been previously shown that circumferential stiffness tends to be maintained well in ANG II-induced hypertension (4) as well as in many cases of genetic defects, including those that affect smooth muscle contractile proteins (3). Tables A1 and A2 in the appendix provide detailed values of Pearson’s correlation coefficient and associated P values for the many possible correlations.

Fig. 6.

Individual values of passive geometric (deformed wall thickness and outer diameter and axial stretch) and mechanical (biaxial wall stress, material stiffness, and energy storage) metrics plotted versus the individual maximum percent change (Δ) in outer diameter in response to 80 mM KCl. Symbols correspond to ANG II-infused (duration of 13–28 days) specimens delineated based on their ability to vasoconstrict (see Fig. 4) as follows: nearly normal (dark gray circles, n = 3), compromised (light gray squares, n = 16), and impaired (white inverted triangles, n = 4). Best-fit linear regression lines are plotted with solid lines to indicate a strong correlation (Pearson’s linear correlation coefficient |ρ| > 0.5 and P < 0.05) or a dashed line (axial stiffness only) to indicate a moderate correlation (ρ ≈ 0.5 and P < 0.1), all based on ANG II-infused samples only. See Table A1 in the appendix for a complete listing of correlations. Noninfused (control, black circles, n = 5) data are overlaid for comparison. All values were calculated at a common pressure of 100 mmHg but individual values of in vivo axial stretch.

Potential impacts on wall stress.

Increases in wall stress stimulate matrix turnover via mechanosensitive responses by smooth muscle cells and (myo)fibroblasts (9, 20). Mean circumferential wall stress can be computed easily using the aforementioned Laplace equation (Eq. 1), where inner radius (a) and wall thickness (h) depend on both distending pressure (P) and smooth muscle contractility. Because increased contractility acts against pressure-induced distensions to reduce luminal diameter and increase wall thickness, the associated mean circumferential Cauchy stress is reduced during smooth muscle contraction relative to relaxation at an equal pressure. We calculated directly, from online measurements, the mean biaxial wall stresses at each of the three fixed pressures (70, 90, and 110 mmHg) used during the vasoconstriction tests at individual mouse-specific axial stretches. Table 3 shows percent reductions in circumferential stress due to high K⁺-induced smooth muscle contraction, with vasoconstriction ranging from 20% to 45% for noninfused control mice to 14–36% for ANG II-infused mice, all for isobaric loading at pressures ranging from 70 to 110 mmHg. Note that samples with impaired vasoconstriction reduced circumferential stress by only 10% in response to high K⁺ at the lowest pressure (70 mmHg) and only reduced axial stress by 5%; they were unable to contract the distended wall against the higher pressures (90 and 110 mmHg). Of course, baseline values of passive circumferential stress were lowest in specimens with an impaired ability to vasoconstrict because of the associated thickening of the wall and reduced axial stretch in hypertension. Reductions in axial stress due to smooth muscle contraction were less than those in circumferential stress but still were ~10–13% in the control group and 8–14% in the ANG II-infused group. Albeit of lower magnitude, these nonzero reductions in axial stress remind us of the potential importance of biaxial contractility (15, 30).

Table 3.

Active behavior of noninfused and ANG II-infused specimens delineated based on their ability to vasoconstrict (see Fig. 4) as follows: near normal, compromised, and impaired

	Circumferential Stress, kPa			Axial Stress, kPa			Circumferential Stretch			Reduction in Diameter
	Relaxed	Contracted	Reduction	Relaxed	Contracted	Reduction	Relaxed	Contraction	Reduction	µm	%
70 mmHg
Control (n = 5)	100 ± 2	55 ± 2	45 ± 2	142 ± 12	128 ± 9	14 ± 2	1.48 ± 0.02	1.12 ± 0.02	0.36 ± 0.02	259 ± 15	22.0 ± 1.1
ANG II infused
Normal (n = 3)	66 ± 3	33 ± 3	32 ± 2	167 ± 25	151 ± 22	16 ± 4	1.31 ± 0.03	0.96 ± 0.05	0.35 ± 0.02	265 ± 22	22.9 ± 1.7
Compromised (n = 16)	56 ± 3	36 ± 2	20 ± 1	101 ± 7	92 ± 7	9 ± 1	1.35 ± 0.03	1.10 ± 0.02	0.25 ± 0.01	181 ± 5	15.0 ± 0.4
Impaired (n = 4)	29 ± 3	26 ± 2	3 ± 2	56 ± 6	52 ± 7	3 ± 2	1.12 ± 0.07	1.06 ± 0.05	0.06 ± 0.03	43 ± 19	3.4 ± 1.6
90 mmHg
Control (n = 5)	171 ± 4	99 ± 4	73 ± 3	143 ± 14	125 ± 11	17 ± 4	1.71 ± 0.02	1.31 ± 0.02	0.40 ± 0.02	289 ± 13	21.5 ± 0.9
ANG II Infused
Normal (n = 3)	113 ± 6	60 ± 4	53 ± 4	179 ± 28	155 ± 22	24 ± 8	1.50 ± 0.04	1.11 ± 0.05	0.39 ± 0.02	299 ± 23	22.8 ± 1.4
Compromised (n = 16)	94 ± 6	63 ± 4	31 ± 2	114 ± 8	100 ± 7	14 ± 2	1.52 ± 0.03	1.27 ± 0.03	0.26 ± 0.01	192 ± 7	14.4 ± 0.5
110 mmHg
Control (n = 5)	244 ± 6	195 ± 8	49 ± 3	149 ± 18	129 ± 14	19 ± 5	1.84 ± 0.02	1.64 ± 0.03	0.19 ± 0.02	141 ± 11	9.8 ± 0.8
ANG II Infused
Normal (n = 3)	180 ± 10	120 ± 4	60 ± 6	206 ± 34	168 ± 24	38 ± 10	1.69 ± 0.02	1.40 ± 0.03	0.29 ± 0.02	231 ± 18	15.7 ± 1.1
Compromised (n = 16)	135 ± 8	109 ± 8	27 ± 2	127 ± 10	108 ± 8	19 ± 3	1.64 ± 0.04	1.47 ± 0.04	0.16 ± 0.01	129 ± 9	9.1 ± 0.6

Values are means ± SE. Relaxed values refer to prestimulus measurements; contracted values refer to poststimulus measurements (80 mM KCl for 15 min). Reduction is the negative of the (contracted- relaxed) value. Mean circumferential and axial stresses were calculated using standard formulas (Eqs. 1 and 2); circumferential stretch is the current value of the midwall radius divided by the unloaded value. The percent reduction in outer diameter (OD) can be calculated as follows: 100 × (OD prestimulus − OD poststimulus)/(OD prestimulus). Note the absence of “impaired” rows at 90 and 110 mmHg since some samples were unable to vasoconstrict against the higher pressures. Note, too, that the ability of an ANG II-infused vessel to reduce diameter or biaxial stress depended on the loading conditions.

DISCUSSION

It has long been thought that normal developmental processes lead to target values of flow-induced wall shear stress and pressure-induced mean circumferential stress in arteries (12, 24, 42). Associated “mechanical homeostasis” also appears to drive many responses in maturity, including adaptive responses to modest changes in hemodynamics (11) and to some extent pathological responses such as aneurysmal enlargement in the absence of comorbidities such as atherosclerosis or thrombosis (41). Indeed, there is extensive evidence that arteries attempt to remodel in hypertension so as to restore the pressure-induced circumferential stress (18, 21, 24) or material stiffness (4, 5) toward normal values. Such responses result, in part, from intramural cells mechanosensing the mechanical state of the extracellular matrix (9, 20) and then mechanoregulating changes to wall composition and organization (32, 38). It has been shown, for example, that arterial remodeling is attenuated in ANG II-induced hypertension in α₁-integrin knockout mice (29), presumably due to a partial loss of mechanomediated remodeling. Despite these many reports, most studies of large vessel remodeling have not attempted to relate changes in geometry or passive wall properties will smooth muscle contractility.

There were three primary findings of this study. First, ANG II-induced remodeling of the thoracic aorta was independent of the degree and duration of pressure elevation measured by tail cuff, at least over the ranges studied. This observation suggests that intramural cells respond first to any sustained mechanostimulus (probably when above a threshold) and perhaps only subsequently fine tune this response. This possibility is consistent with longer-term findings for the carotid artery in another mouse model of induced hypertension (13). Second, the hypertensive remodeling was characterized by marked mouse-to-mouse differences, with 3 of 23 mice having mechanoadapted DTAs but 20 of 23 mice exhibiting a dramatic range of maladaptation over the 2- to 4-wk study. Given these percentages, it is not surprising that prior studies using smaller sample sizes (often n = 5–7) have reported maladaptive responses alone (4, 43). Third, the extent of overall remodeling correlated strongly and inversely with the ability of the aorta to vasoconstrict in response to 80 mM KCl, with increasingly maladapted vessels (evidenced by geometric, microstructural, and passive property abnormalities) having an increasingly limited ability to vasoconstrict. This is, to our knowledge, the first strong correlation identified between altered passive and active properties in induced hypertension in mice.

Indeed, although most studies in rodents have similarly reported increased matrix deposition with thickening and structural stiffening of the aortic wall in hypertension (e.g., Refs. 31 and 39), a number of studies have conversely reported an increase in active smooth muscle force generation. For example, Gallardo-Ortiz et al. (19) reported that contractions to 80 mM KCl were stronger in rings of the thoracic aorta when excised from rats subjected to low-dose ANG II infusion (200 ng·kg⁻¹·min⁻¹ for 2 wk) than when excised from noninfused control mice. There are, however, two key differences between their study and our own: species (rat vs. mouse) and method of testing (uniaxial versus biaxial). Yet, Fransen et al. (17) also found that KCl-induced contractions were more forceful in rings of the thoracic aorta of wild-type mice subjected to high-dose ANG II infusion (1,000 ng·kg⁻¹·min⁻¹ for 4 wk) than when excised from noninfused mice. Similarly, phenylephrine-induced contractions were increased in aortic rings excised from both rats (19) and mice (1) after ANG II infusion. Hence, known species differences may not be the issue here. Two concerns common to these three studies, however, are that the preload was the same on the infused and noninfused samples and the measurement was of force of contraction, not the associated stress generation. In other words, these studies did not account for potential differences in loading or wall thickness due to hypertensive remodeling. In contrast, we considered up to nine different biaxial preloading states, with responses depending on the preload (Table 3), and we used normalized metrics (normalized outer diameter and wall stress) since strain-like and stress-like quantities are always preferred for material characterization. Importantly, Fitch et al. (16) compared contractility using the same type of ring myography that was used in these three other studies (1, 17, 19), but they reported the active generation of aortic wall stress (in mN/mm²), not force (in mN). Inferred in this way, contractile capacity was statistically lower in DTAs excised from mice that were infused with ANG II (500 ng·kg⁻¹·min⁻¹ for 4 wk), consistent with our findings at a higher dose of ANG II (1,000 ng·kg⁻¹·min⁻¹ for 2 to 4 wk). Although they reported the active generation of force, not stress, Prasad et al. (31) also found a reduced contractile response by murine aortic rings exposed to KCl after high-dose ANG II infusion (1,250 ng·kg⁻¹·min⁻¹ for 2 wk), although the contractile response was higher when stimulated with phenylephrine. Although one should always compare current with prior results, comparisons across studies should also address the approach of mechanical testing, not just the dose of ANG II or the vasostimulant used during the in vitro assessment. That is, notwithstanding the importance of using normalized mechanical metrics (e.g., stress rather than force), we suggest further that the present biaxial results are more relevant to the in vivo condition, and thus novel, since we maintained the vessels in vitro in a native cylindrical geometry under physiological pressures and axial stretches and we did not purposely denude the endothelium. Although uniaxial ring tests are much easier to perform, it is preferable for both mechanical and physiological reasons to test vessels under more natural conditions (10, 22). The interested reader is referred to a direct comparison of uniaxial ring, biaxially isometric, and biaxial isobaric-isometric testing, which shows particular advantages of biaxial tests (8), not the least of which is appropriate control of the axial stretch, which changes dramatically in hypertension and affects biaxial stress, stiffness, energy storage, and contractility.

We also found that aortic maladaptation was evidenced not only by an excessive thickening of the wall but also by a marked encroachment upon the lumen. Based on the mean increase in ambulatory blood pressure (γ = 167/121 = 1.38) in the presence of an assumed unchanged cardiac output (ε ≈ 1), mean circumferential wall stress and wall shear stress would theoretically be maintained at or restored to near normal values if wall thickness increased by 38% while inner radius remained nearly the same (23). The 3 of 23 ANG II-treated DTAs having a near normal ability to vasoconstrict exhibited a 29% increase in wall thickness (49 vs. 38 µm at baseline), which was perhaps beginning to approach the optimal 38% increase, and a 12% increase in inner radius (831 vs. 740 µm at baseline), perhaps consistent with the increased distending pressure and yet incomplete adaptation. The increased wall thickness resulted from increases in both medial and adventitial thickness (67% and 34% increases, respectively, in unloaded cross-sectional area), with the additional adventitial collagen fibers tending to be thinner (11% and 25% increases in the area fraction of yellow and green fiber relative to baseline, respectively). Importantly, however, there were no statistically significant differences in axial stretch, wall stress, stiffness, energy storage, or contractile capacity in these three vessels, consistent with a favorable mechanoadaptation.

In stark contrast, 20 of 23 ANG II-treated DTAs showed a dramatic range of maladaptations, evidenced by 1.1- to 3.4-fold increases in wall thickness, a generally marked change in inner radius (ranging from 581 to 848 µm relative to 740 µm at baseline), and a marked change in elastic energy storage (ranging from 11 to 106 kPa relative to 78 kPa at baseline). Largely due to the excessive thickening, circumferential stress was also markedly different than baseline (ranging from 101 to 498 kPa at elevated P_sys relative to 312 kPa at normal P_sys) and the wall shear stress was likely elevated due to the generally reduced caliber. Interestingly, these maladaptations occurred despite a near maintenance of circumferential material stiffness, which suggests a preserved ability of the intramural cells to mechanosense and mechanoregulate a key extracellular matrix property (2). Proper mechanosensing and mechanoregulation of matrix requires competent actomyosin activity and associated intracellular signaling pathways (26). MTC staining and immunostaining revealed mild reductions in (not complete losses of) smooth muscle α-actin and myosin heavy chain with maladaptation despite the apparent mechanobiological control of circumferential matrix stiffness. Nevertheless, the ability to vasoconstrict under isobaric-axially isometric conditions was much lower in maladapted hypertensive vessels than in either the normotensive or the near mechanoadapted hypertensive vessels. Clearly, therefore, functional studies are critical in assessing potential consequences of molecular or histological findings. There is, however, a continuing need to better understand how actomyosin activity is used by cells for microstructural (e.g., sensing) versus macroscale (e.g., vasoconstriction) regulation. There is also a need to determine whether a loss of contractile capacity contributes to or is a consequence of the maladaptive remodeling or both. Although one would expect a modulation of smooth muscle cells from a mature contractile to a more synthetic phenotype in hypertension, which should reduce the ability of the vessel to constrict while increasing matrix synthesis, it is possible that early defects in contractility prevent smooth muscle cells from carrying part of the pressure-induced load (Table 3), thus resulting in higher initial wall stress and a greater thickening response. The latter possibility is consistent with a recent hypothesis (5).

Interestingly, the luminal encroachment that associated with maladaptation (e.g., inner radius of ~613 µm in the impaired vasoconstriction samples relative to 740 µm in controls; Table 2) occurred in the absence of neointimal formation (Fig. 5) or an in vitro ability of the vessels to vasoconstrict to 80 mM KCl (Fig. 4). It is possible that the smooth muscle cells could have contracted in response to other agonists, but this was not evaluated. It may also be that a thicker and stiffer adventitia, with a higher load-bearing capability than the media (4), simply caused the medial thickening to proceed inward, not unlike in stenotic neotissue development in bilayered tissue engineered vascular grafts having a stiff outer sheath (37). That is, endothelial function [not measured, but often dysfunctional in hypertension (27, 35)] may not have played much of a mechanical role in the maladaptation, although increased nitric oxide [which is anti-inflammatory but often limited in ANG II infusion studies (36)] may have helped to limit neointimal development. In contrast, the present finding of adventitial fibrosis, with increased myofibroblast density, in the maladapted vessels is consistent with recent reports showing that mechanoadaptations can be compromised by excessive inflammatory activity in the adventitia. It has been shown in male mice, for example, that CD45⁺ cells contribute significantly to the structural stiffening of the DTA in induced hypertension (4, 43). Our assessment of CD45⁺ cells similarly revealed an increase with ANG II infusion (not shown), although without a clear correlation with diminished smooth muscle contractile capability. This latter observation appears consistent with our prior report that inflammation appears to follow rather than precede increased wall stress and thus mechanostimulated matrix turnover in response to early increases in blood pressure (5).

In summary, our results are consistent with the longstanding thought that aortic wall remodeling in hypertension depends strongly on intramural wall stress but suggest further that the (in)ability of a vessel to vasoconstrict may play a role in dictating the level of wall stress and thus the remodeling response. Importantly, mouse-to-mouse differences in smooth muscle contractile capacity, either intrinsic or in response to infused ANG II, correlated strongly with the biomechanical phenotype, which in some cases was characterized by a grossly maladaptive thickening and thus structural stiffening of the wall. Hence, although it has long been known that smooth muscle contractility plays a critical role in hypertension, with increased microvascular tone adversely increasing systemic vascular resistance (R) (7, 40), which necessitates an increase in driving pressure (ΔP) to maintain the volumetric flow rate (Q) (with ΔP = QR), we submit that there is also a need to consider smooth muscle function in large vessels wherein most of the compliance of the vasculature resides (34), with assessments based on a range of vasostimulants, not just KCl. Increased vasoconstriction of large vessels could help reduce the maladaptive structural stiffening that ultimately increases pulse wave velocity and thereby augments central pulse pressure. This dichotomy (detrimental increased smooth muscle contraction in microvessels versus advantageous increased contraction in large vessels) may be similar to competitions between local mechanobiology and global hemodynamics wherein thickening of the wall can favorably reduce local wall stress but adversely increase structural stiffening (25, 26). Vascular health thus appears to rely on delicate balances. From the perspective of smooth muscle contractility, the present findings suggest that pharmacologic interventions that relax microvessels but not macrovessels may be useful in treating some types of hypertension.

Finally, as we move into the era of precision (personalized) medicine, the phenotypic diversity exhibited within our otherwise similar cohort (i.e., same genotype, sex, adult age, diet, exercise, insult, and so forth) emphasizes the challenge of understanding individual responses. Of course, our assessment of vasoconstriction capacity and correlation with the degree of aortic wall remodeling was necessarily based on vessels that had prior long-term exposure to ANG II. Hence, we could not delineate whether there was a de novo mouse-to-mouse difference in smooth muscle contractile capacity (assessed by KCl) or whether there was a differential response to the infused ANG II. Regardless, the importance of mouse-specific contractile capacity remains. Indeed, we recently reported that mice with a germline mutation in the gene that encodes smooth muscle myosin heavy chain (Myh11) exhibited a surprisingly mild phenotype unless rendered hypertensive (with high NaCl and N-nitro-l-arginine methyl ester). Approximately 25% of the hypertensive mutants died from an aortic dissection, whereas ~20% of the survivors exhibited an intramural delamination; it was not clear, however, why most mice exhibited no phenotype over the period of study (3). Again, there is a need to understand better mouse-to-mouse differences even within well-controlled populations. Another key issue is also not whether or not a phenotype will manifest but at what time it might manifest. That is, we should begin to think more about biological, not just chronological, time. Mouse-to-mouse (and similarly subject-to-subject) variability, including temporal manifestation, needs to be understood better. Toward this end, it will become increasingly important not to exclude apparent outliers in small cohorts since their inclusion could provide greater insights into the disease progression and variability across individuals (3). Given that very different genetic backgrounds (including race and ethnicity), sexes, ages, diets, levels of exercise, and states of emotional stress define a normal clinical population, the challenge is considerable. With regard to the present study, that the extent of aortic remodeling, that is, loss of mechanical functionality revealed as lower elastic energy storage and increased structural stiffening caused primarily by increased wall thickness, correlated inversely with mouse-specific smooth muscle cell contractile capacity, intrinsic phenotypic diversity in smooth muscle (and other vascular) cells may similarly be responsible for part of the variability seen in the human patient population and thus should be increasingly scrutinized.

GRANTS

This work was supported, in part, by National Institutes of Health Grants R01 HL-105297, U01 HL-116323, R21 EB-020968, and R03 EB-021430.

DISCLOSURES

No conflicts of interest, financial or otherwise, are declared by the authors.

AUTHOR CONTRIBUTIONS

A.K. and J.D.H. conceived and designed research; A.K. performed experiments; A.K. and J.D.H. analyzed data; A.K. and J.D.H. interpreted results of experiments; A.K. prepared figures; J.D.H. drafted manuscript; A.K. and J.D.H. edited and revised manuscript; A.K. and J.D.H. approved final version of manuscript.

APPENDIX

Tables A1 and A2 show detailed values of Pearson’s correlation coefficient and the associated P values for the many possible correlations.

Table A1.

Pearson's linear correlation coefficient was calculated between individual geometric and biaxial passive metrics and one active metric: the value of maximum percent change in outer diameter upon stimulation with 80 mM KCl

	Max. ΔOuter Diameter, %	Loaded Thickness	Loaded Outer Diameter	In Vivo Axial Stretch, ${\text{λ}}_z^{\text{iv}}$	Loaded Circ. Stretch, λθ	Stored Energy	Axial Stress, σz	Circ. Stress, σθ	Axial Stiffness, Czzzz	Circ. Stiffness, Cθθθθ	Unloaded Thickness	Unloaded Outer Diameter	Energy Dissipation Ratio	Systolic Pressure	Infusion Duration	Age
Max. Δouter diameter, %		−0.7*	0.6*	0.5*	0.6*	0.7*	0.6*	0.7*	0.5	0.3	−0.6*	−0.3	−0.4	0.1	0.3	−0.4
Loaded thickness	−0.7*		−0.5	−0.8*	−0.8*	−0.9*	−0.9*	−0.9*	−0.8*	−0.4	0.9*	0.6*	0.3	0.2	−0.1	0.2
Loaded outer diameter	0.6*	−0.5		0.6*	0.3	0.7*	0.6*	0.7*	0.3	0.5*	−0.5	0.3	−0.1	0.2	−0.0	−0.3
In vivo axial Stretch, ${\text{λ}}_z^{\text{iv}}$	0.5*	−0.8*	0.6*		0.6*	0.9*	0.9*	0.8*	0.6*	0.5*	−0.7*	−0.3	−0.2	0.0	0.1	−0.2
Loaded circ. stretch, λθ	0.6*	−0.8*	0.3	0.6*		0.8*	0.8*	0.8*	0.8*	0.3	−0.6*	−0.8*	−0.1	−0.4	−0.0	−0.2
Stored energy	0.7*	−0.9*	0.7*	0.9*	0.8*		1.0*	1.0*	0.8*	0.6*	−0.8*	−0.4	−0.1	−0.1	0.1	−0.3
Axial stress, σz	0.6*	−0.9*	0.6*	0.9*	0.8*	1.0*		1.0*	0.9*	0.5*	−0.9*	−0.5	−0.2	−0.2	0.1	−0.2
Circ. stress, σθ	0.7*	−0.9*	0.7*	0.8*	0.8*	1.0*	1.0*		0.8*	0.6*	−0.9*	−0.4	−0.1	−0.1	0.0	−0.3
Axial stiffness, Czzzz	0.5	−0.8*	0.3	0.6*	0.8*	0.8*	0.9*	0.8*		0.4	−0.7*	−0.7*	0.0	−0.4	0.1	−0.2
Circ. stiffness, Cθθθθ	0.3	−0.4	0.5*	0.5*	0.3	0.6*	0.5*	0.6*	0.4		−0.4	−0.1	0.3	0.3	0.1	−0.1
Unloaded thickness	−0.6*	0.9*	−0.5	−0.7*	−0.6*	−0.8*	−0.9*	−0.9*	−0.7*	−0.4		0.4	0.3	0.2	−0.1	0.2
Unloaded outer diameter	−0.3	0.6*	0.3	−0.3	−0.8*	−0.4	−0.5	−0.4	−0.7*	−0.1	0.4		0.1	0.5	−0.1	0.0
Energy dissipation ratio	−0.4	0.3	−0.1	−0.2	−0.1	−0.1	−0.2	−0.1	0.0	0.3	0.3	0.1		−0.0	−0.3	−0.0
Systolic pressure	0.1	0.2	0.2	0.0	−0.4	−0.1	−0.2	−0.1	−0.4	0.3	0.2	0.5	−0.0		0.5	−0.1
Infusion duration	0.3	−0.1	−0.0	0.1	−0.0	0.1	0.1	0.0	0.1	0.1	−0.1	−0.1	−0.3	0.5		0.4
Age	−0.4	0.2	−0.3	−0.2	−0.2	−0.3	−0.2	−0.3	−0.2	−0.1	0.2	0.0	−0.0	−0.1	0.4

Unless specified as unloaded, values were calculated at a common pressure of 100 mmHg but individual in vivo axial stretches; all values were rounded to one decimal point for ease of tabulation. Additional correlation coefficients were computed for three experimental variables: mouse-specific systolic blood pressure, duration of infusion, and mouse age. Max., maximum; circ., circumferential.

^*Strong linear correlations (|ρ| ≥ 0.5, where ρ is Pearson's linear correlation coefficient) are indicated. Note, too, that these indicated values also correspond to associated P ≤ 0.01 (Table A2).

Table A2.

Similar to Table A1 except for associated P values

	Max. ΔOuter Diameter, %	Loaded Thickness	Loaded Outer Diameter	In Vivo Axial Stretch, ${\text{λ}}_z^{\text{iv}}$	Loaded Circ. Stretch, λθ	Stored Energy	Axial Stress, σz	Circ. Stress, σθ	Axial Stiffness, Czzzz	Circ. Stiffness, Cθθθθ	Unloaded Thickness	Unloaded Outer Diameter	Energy Dissipation Ratio	Systolic Pressure	Infusion Duration	Age
Max. Δouter diameter, %		0.0*	0.01*	0.01*	0.0*	0.0*	0.0*	0.0*	0.02	0.13	0.0*	0.17	0.08	0.71	0.25	0.09
Loaded thickness	0.0*		0.03	0.0*	0.0*	0.0*	0.0*	0.0*	0.0*	0.07	0.0*	0.0*	0.18	0.3	0.65	0.3
Loaded outer diameter	0.01*	0.03		0.0*	0.23	0.0*	0.0*	0.0*	0.13	0.01*	0.03	0.2	0.74	0.49	0.85	0.12
In vivo axial stretch, ${\text{λ}}_z^{\text{iv}}$	0.01*	0.0*	0.0*		0.0*	0.0*	0.0*	0.0*	0.0*	0.01*	0.0*	0.23	0.3	0.99	0.68	0.35
Loaded circ. stretch, λθ	0.0*	0.0*	0.23	0.0*		0.0*	0.0*	0.0	0.0*	0.15	0.0*	0.0*	0.74	0.1	0.9	0.27
Stored energy	0.0*	0.0*	0.0*	0.0*	0.0*		0.0*	0.0*	0.0*	0.0*	0.0*	0.04	0.49	0.56	0.64	0.25
Axial stress, σz	0.0*	0.0*	0.0*	0.0*	0.0*	0.0*		0.0*	0.0*	0.01*	0.0*	0.02	0.39	0.49	0.58	0.27
Circ. stress, σθ	0.0*	0.0*	0.0*	0.0*	0.0*	0.0*	0.0*		0.0*	0.0*	0.0*	0.04	0.62	0.54	0.85	0.2
Axial stiffness, Czzzz	0.02	0.0*	0.13	0.0*	0.0*	0.0*	0.0*	0.0*		0.08	0.0*	0.0*	0.98	0.11	0.81	0.44
Circ. stiffness, Cθθθθ	0.13	0.07	0.01*	0.01*	0.15	0.0*	0.01*	0.0*	0.08		0.04	0.81	0.19	0.2	0.61	0.63
Unloaded thickness	0.0*	0.0*	0.03	0.0*	0.0*	0.0*	0.0*	0.0*	0.0*	0.04		0.04	0.25	0.48	0.6	0.44
Unloaded outer diameter	0.17	0.0*	0.2	0.23	0.0*	0.04	0.02	0.04	0.0*	0.81	0.04		0.75	0.03	0.78	0.88
Energy dissipation ratio	0.08	0.18	0.74	0.3	0.73	0.49	0.39	0.62	0.98	0.19	0.25	0.75		0.93	0.12	0.88
Systolic pressure	0.71	0.3	0.49	0.99	0.1	0.56	0.49	0.54	0.11	0.2	0.48	0.03	0.93		0.04	0.54
Infusion duration	0.25	0.65	0.85	0.68	0.9	0.64	0.58	0.85	0.81	0.61	0.6	0.78	0.12	0.04		0.04
Age	0.09	0.3	0.12	0.35	0.27	0.25	0.27	0.2	0.44	0.63	0.44	0.88	0.88	0.54	0.04

All values were rounded to 2 decimal points for ease of tabulation. Max., maximum; circ., circumferential.

^*Values corresponding to a strong linear correlation, with Pearson correlation coefficient |ρ| ≥ 0.5.