A novel set‐up for the ex vivo analysis of mechanical properties of mouse aortic segments stretched at physiological pressure and frequency

Abstract

Key points

• Cyclic stretch is known to alter intracellular pathways involved in vessel tone regulation.

• We developed a novel set‐up that allows straightforward characterization of the biomechanical properties of the mouse aorta while stretched at a physiological heart rate (600 beats min^–1).

• Active vessel tone was shown to have surprisingly large effects on isobaric stiffness.

• The effect of structural vessel wall alterations was confirmed using a genetic mouse model.

• This set‐up will contribute to a better understanding of how active vessel wall components and mechanical stimuli such as stretch frequency and amplitude regulate aortic mechanics.

Abstract

Cyclic stretch is a major contributor to vascular function. However, isolated mouse aortas are frequently studied at low stretch frequency or even in isometric conditions. Pacing experiments in rodents and humans show that arterial compliance is stretch frequency dependent. The Rodent Oscillatory Tension Set‐up to study Arterial Compliance is an in‐house developed organ bath set‐up that clamps aortic segments to imposed preloads at physiological rates up to 600 beats min^–1. The technique enables us to derive pressure–diameter loops and assess biomechanical properties of the segment. To validate the applicability of this set‐up we aimed to confirm the effects of distension pressure and vascular smooth muscle tone on arterial stiffness. At physiological stretch frequency (10 Hz), the Peterson modulus (E _P; 293 (10) mmHg) for wild‐type mouse aorta increased 22% upon a rise in pressure from 80–120 mmHg to 100–140 mmHg, while, at normal pressure, E _P increased 80% upon maximal contraction of the vascular smooth muscle cells. We further validated the method using a mouse model with a mutation in the fibrillin‐1 gene and an endothelial nitric oxide synthase knock‐out model. Both models are known to have increased arterial stiffness, and this was confirmed using the set‐up. To our knowledge, this is the first set‐up that facilitates the study of biomechanical properties of mouse aortic segments at physiological stretch frequency and pressure. We believe that this set‐up can contribute to a better understanding of how cyclic stretch frequency, amplitude and active vessel wall components influence arterial stiffening.

Key points

• Cyclic stretch is known to alter intracellular pathways involved in vessel tone regulation.

• We developed a novel set‐up that allows straightforward characterization of the biomechanical properties of the mouse aorta while stretched at a physiological heart rate (600 beats min^–1).

• Active vessel tone was shown to have surprisingly large effects on isobaric stiffness.

• The effect of structural vessel wall alterations was confirmed using a genetic mouse model.

• This set‐up will contribute to a better understanding of how active vessel wall components and mechanical stimuli such as stretch frequency and amplitude regulate aortic mechanics.

Abbreviations

CV

cardiovascular

eNOS

endothelial nitric oxide synthase

l‐NAME

N ^ω‐nitro‐l‐arginine methyl ester

NO

nitric oxide

PE

phenylephrine

ROTSAC

Rodent Oscillatory Tension Set‐up to study Arterial Compliance

VSMC

vascular smooth muscle cell

Introduction

Progressive large artery stiffening is the predominant cause of increased pulse pressure, a marker of cardiovascular (CV) risk in the general population (Benetos et al. 1997) and a predictor of CV events (Mitchell et al. 1997). Furthermore, it reduces myocardial perfusion efficiency, increases left ventricular afterload and elicits mechanical stress on capillaries, potentially damaging the capillary wall of strongly perfused organs such as the heart, brain and kidneys (Safar et al. 2012). It is generally assumed that arterial stiffness is an adaptation mechanism of the aortic wall to increased distending pressures. However, the REASON study, published in 2009, states that, in hypertensive patients, high arterial stiffness predicts a poor response to antihypertensive treatment (Protogerou et al. 2009). In addition, several epidemiological studies report that arterial stiffness precedes hypertension in the elderly (Najjar et al. 2008; Kaess et al. 2012). This temporal relationship was recently confirmed in a mouse model of diet‐induced obesity; i.e. arterial stiffness preceded hypertension and target organ damage in this model (Weisbrod et al. 2013). These observations suggest a role for arterial stiffness as a therapeutic target to treat CV disease.

In recent years, there is emerging evidence that it is not only passive components that determine arterial compliance. Indeed, vascular smooth muscle cell (VSMC) stiffness and active vessel wall components (i.e. nitric oxide (NO) bioavailability and VSMC tonus) also affect arterial compliance (Fitch et al. 2001; Kinlay et al. 2001; Bellien et al. 2010; Sehgel et al. 2013). Although the interest in these pathways as potential therapeutic targets has grown, their definite role in arterial stiffening during disease processes is still incompletely understood, increasing the need for further research on active vessel wall components in arterial stiffening. Several in vivo techniques to measure arterial stiffness are available (Hartley et al. 1997; Herold et al. 2009; Di Lascio et al. 2014; Leloup et al. 2014), but in order to analyse the underlying molecular pathways, ex vivo analysis with manipulation of the experimental environment is required. Ex vivo perfusion of artery segments is frequently used to assess biomechanical parameters (Huang et al. 2008; Le et al. 2011), but the lack of cyclic stretch may alter endothelial NO synthase (eNOS) activity, especially since aortic endothelial cells are sensitive to cyclic stretch amplitude (Peng et al. 2003) and heart rate is a blood pressure‐independent determinant of aortic stiffness (Millasseau et al. 2005; Tan et al. 2012). In addition, arteries are known to show viscoelastic behaviour as indicated by the hysteresis phenomenon in the pressure–diameter plot (Boutouyrie et al. 1997) and this depends on the frequency of stretching. Indeed, when stretch frequency is low, the pressure–diameter relationship is linear and hysteresis is absent (Santelices et al. 2007). The viscosity of the artery is thought to be a major contributor to the pulse‐smoothening capacity of the aorta, indicating that the study of VSMC tone and aortic mechanics under isometric conditions in the commonly used wire or perfusion myograph set‐ups may be biased. Pulsatile perfusion can be applied but, given the high heart rate and small size of the mouse, this is generally applied to larger animal models such as the rat (Boutouyrie et al. 1997) and sheep (Bia et al. 2005). A set‐up that exposes mouse aortic segments to pulsatile perfusion has been reported but only with subphysiological stretch rate (20 beats min^–1) (Santelices et al. 2007).

Therefore, we aimed to develop a novel ex vivo approach to study elastic behaviour of aortic segments while stretched at physiological frequencies and amplitudes. The Rodent Oscillatory Tension Set‐up to study Arterial Compliance (ROTSAC) is an organ bath set‐up with the aortic segment connected to a force–length transducer that allows clamping the segments at different preloads. A stretch protocol was applied with physiological frequency (10 Hz) and amplitude (40 mmHg). We aimed to (i) estimate the reproducibility of the set‐up, (ii) confirm the ex vivo effect of increased distension pressure and VMSC tone on stiffness parameters, (iii) assess aortic stiffness of isolated aortic segments from two genetic mouse models known to have increased aortic stiffness, and (iv) test whether physical integrity of the isolated aortic segment is morphologically and mechanically maintained when high frequency oscillation is applied.

Methods

Ethical approval

The studies were approved by the Ethical Committee of the University of Antwerp, and all experiments were performed conform to the Guide for the Care and Use of Laboratory Animals published by the US National Institutes of Health (NIH Publication No. 85–23, revised 1996).

Animals

All animals (C57Bl/6J background) were bred and housed in the animal facility of the University of Antwerp in standard cages with 12 h–12 h light–dark cycles with free access to regular chow and tap water. The length and diameter calibration method was validated using eight aortic segments from four male WT mice of mean (SD) weight 28.0 (1.2) g. For the other experiments, five male WT mice of 22.5 (2.0) g, five male eNOS^−/− mice of 30.6 (3.4) g and four male ApoE^−/−Fbn1^C1039G+/− (Fbn‐1‐mut) mice of 32.7 (0.8) g) were used. Animals were killed by perforating the diaphragm while under anaesthesia (sodium pentobarbital (Sanofi, Belgium), 75 mg kg⁻¹, i.p.). The thoracic aorta was carefully removed and stripped of adherent tissue. Starting at the diaphragm, the aorta was cut into six segments of 2 mm width. The third and fourth segments were immersed in Krebs–Ringer solution (37°C, 95% O₂–5% CO₂, pH 7.4) containing (in mm): NaCl 118, KCl 4.7, CaCl₂ 2.5, KH₂PO₄ 1.2, MgSO₄ 1.2, NaHCO₃ 25, CaEDTA 0.025 and glucose 11.1. To avoid any vasomotor interference due to prostanoids, 10 μm indomethacin (Federa, Belgium) was present in all experiments.

ROTSAC

Aortic segments were mounted between two parallel wire hooks in 8 ml organ baths. Force and displacement of the upper hook were controlled and measured with a force–length transducer described earlier (Brutsaert et al. 1971; Brutsaert & Claes, 1974). In short, the upper hook was connected to the aluminium lever of the force–length transducer. This lever was connected to a coil suspended in a strong field of a permanent magnet. The system was controlled by a current source. When current was passed through this coil, a force was developed. The displacement of the lever was measured by means of a photo‐electric system. The transducer was connected to a data acquisition system (Powerlab 8/30 and LabChart 7, ADInstruments, Oxford, UK) (Fig. 1). Force and displacement were acquired at 1 kHz. To calculate the transmural pressure that would exist in the equilibrated vessel segment with the given wall stress (derived from the distension force) and dimensions, the Laplace relationship was used:

$$P=\frac{F}{l.D}$$

with F the force, l the length and D the diameter of the vessel segment. Force was measured directly by the transducer. The diameter of the vessel segment at a given preload was derived from the displacement of the upper hook, being directly proportional to the inner circumference:

$$D=\frac{2g}{\pi }$$

with g the outer distance between the hooks (to approximate the inner circumference of the vessel segment). Before each experiment, diameter and length were determined at three different preloads (20, 40 and 60 mN) using a binocular and calibrated image software. To correct for the slight decrease in vessel length with increased diameter, the average length per cycle (100 ms) was derived from the diameter–length relationship using linear regression.

Figure 1. Schematic diagram of the ROTSAC .

The aortic segment was mounted between two metal hooks in an organ bath. (1) A current source was used to control the distension force and clamp frequency (10 Hz) of the force–length transducer. Force and displacement were measured by the transducer (2) and acquired at 1 kHz by a PowerLab DAQ device (3). Diameter, length and force were used to calculate the pressure that would exist in an equilibrated vessel segment with the given dimensions and wall stress (4).

The preload was adjusted until the desired calculated diastolic and systolic pressure were attained. The acquired pressure–diameter loops were used for further analysis. Compliance (C) was calculated as follows:

$$C=\frac{\mathrm{\Delta }D}{\mathrm{\Delta }P}$$

with ∆D, the difference between systolic and diastolic diameter and ∆P, the pressure difference (±40 mmHg in the present study). The Peterson modulus of elasticity (E _P) is a frequently used, vessel size‐independent measure of arterial stiffness (Gosling & Budge, 2003) and was calculated as follows:

$$E{}_{\mathrm{P}}=D_0.\frac{\mathrm{\Delta }P}{\mathrm{\Delta }D}$$

with D ₀, the diastolic diameter. During all experiments, the segments were periodically stretched directly after mounting them in the organ bath with a frequency of 10 Hz to mimic the physiological heart rate in mice (600 beats min^–1). VSMCs were stimulated using the α₁‐adrenergic agonist phenylephrine (PE) (Sigma‐Aldrich, Belgium) and N ^ω‐nitro‐l‐arginine methyl ester (l‐NAME) (Sigma‐Aldrich, Belgium) was used to inhibit eNOS.

Histology

Segments, which were mounted in the ROTSAC for 4 h with or without repetitive stretching (10 Hz) at preloads corresponding to normal pressures, were formalin‐fixed for 24 h, dehydrated overnight in 60% isopropanol and embedded in paraffin. Histological analysis was performed on serial cross sections (5 μm) stained with orcein to visualize elastin. The images were acquired with the Universal Grab 6.1. (IDL) software (Exelis, Boulder, CO, USA) using an Olympus BX40 microscope (Tokyo, Japan).

Statistical analysis

All results are expressed as mean (SD) or median (interquartile range) with n representing the number of mice, and analysed using Prism 6.0 (GraphPad Software, La Jolla, CA, USA). Data of the VSMC tone and pressure effect were compared by two‐way ANOVA with Bonferroni multiple comparison post hoc test. The genetic mouse models with arterial stiffness were compared against WT mice using a Mann–Whitney U test. The effect of the 4 h stretch protocol on E _P was evaluated using a Student's paired t test. A 5% level of significance was selected.

Results

Vessel dimensions are accurately derived from the force–length transducer

Effective diameter and length were determined at different preloads using a binocular microscope and compared with the calculated diameter derived from the displacement of the upper hook. Bland–Altman analysis revealed a bias of –10.73 μm on the estimation of the diameter (0.9% of the mean diameter). The 95% limits of agreement were –44.6 μm (3.9% of the mean diameter) and +23.2 μm (2.1% of the mean diameter) (Fig. 2 A and B). When length was determined using the diameter–length relationship, the bias was –4.3 μm (0.3% of the mean length) and 95% limits of agreement were –34.4 μm (2.0% of the mean) and +25.8 μm (0.9% of the mean length) (Fig. 2 C and D). Overall, both estimations of length and diameter were considered sufficiently accurate and this method of estimating vessel dimensions was used to determine the calculated transmural pressure. Figure 3 shows the diameter–time plot (Fig. 3 A) of a compliant and a stiff vessel segment. The hysteresis phenomenon typically seen when stretching vessel segments at sufficiently high frequency is apparent from the resulting pressure–diameter plot (Fig. 3 B).

Figure 2. Bland–Altman analysis of geometrical parameters .

Comparison between effective diameter (D _eff) (A and B) and length (L _eff) (C and D) and the calculated parameters (D _calc and L _calc) derived from the displacement of the needle and the diameter–length relationship, respectively. The dashed lines represent the bias and 95% limits of agreement.

Figure 3. Processing of the data obtained from a compliant and stiff vessel segment .

The graphs show examples of a WT aortic segment, either fully contracted using 1 μm PE and 300 μm l‐NAME (stiff), or in Krebs–Ringer solution (compliant). Pressure (continuous line) and diameter (dashed line) are plotted against time (A). As shown, segments were stretched between 80 and 120 mmHg at a stretch frequency of 10 Hz. When pressure and diameter are plotted against each other, a typical hysteresis loop was obtained (B). Compliant (black) and stiff (grey) vessel segments are shown to illustrate the differences in distension with similar pressure changes (A) and the difference in slope of the pressure–diameter loops (B).

The ROTSAC allows isobaric assessment of mechanical properties of aortic segments with different VSMC tone or at different distension pressures

Because distension pressure and VSMC tone are known modulators of aortic stiffness, we tested whether this set‐up was able to pick up changes in biomechanical properties induced by pharmacological stimulation of VSMCs, both at normal pressure (80–120 mmHg) and higher pressure (100–140 mmHg). Upon pharmacological stimulation of α₁‐adrenergic receptors with 1 μm phenylephrine (PE), diastolic diameter (D₀) significantly decreased and both diameter‐dependent and diameter‐independent parameters changed significantly (Fig. 4). This effect was even more pronounced when the basal production of the relaxing factor NO was inhibited using 300 μm l‐NAME. As expected, increased distending pressure significantly changed the mechanical properties of the aortic segment for all conditions (Fig. 4).

Figure 4. Isobaric vessel parameters under conditions with different VSMC tone and pressure .

Diastolic diameter (D ₀) (A), relative distension (B), compliance (C) and Peterson modulus (E _P) (D) were measured at physiological stretch frequency (10 Hz), either at normal pressure (80–120 mmHg) or at high pressure (100–140 mmHg). KR, Krebs–Ringer; PE, 1 μm phenylephrine; PE+LN, 1 μm phenylephrine + 300 μm l‐NAME. ^* P < 0.05, ^** P < 0.01, ^*** P < 0.001 (effect of VSMC tone, vs KR), ^††† P < 0.001 (effect of pressure, vs. 80‐120 mmHg), two‐way ANOVA with repeated measures for pressure factor, Bonferroni post hoc test (n = 5).

The ROTSAC confirms altered mechanical properties in aortic segments of two genetic mouse models

To test the capacity of the set‐up to pick up differences in arterial stiffness between mouse models with known altered aortic geometry and mechanics, a mouse model with a mutation in the fibrillin‐1 gene (Fbn1) was used (Van der Donckt et al. 2015). The aortic diameter was increased as compared to WT mice (Fig. 5 A). Relative distension (Fig. 5 B) and compliance (Fig. 5 C) were significantly lower and E _P was significantly increased at physiological pressure (Fig. 5 D).

Figure 5. Comparison between WT mice (n = 5) and mice with a heterozygous mutation of the Fbn1 gene (n = 4) .

The aorta was dilated, as indicated by increased diastolic diameter (D ₀) (A). Relative distension (B) and compliance (C) were significantly lower and the Peterson modulus (E _P) was significantly increased (D). Measurements were done at physiological pressure range (80–120 mmHg) and stretch frequency (10 Hz). Line and error bars represent median and interquartile range, respectively. ^* P < 0.05, Mann–Whitney U test.

Next, aortic segments from eNOS^−/− mice – as a second model known to have increased arterial stiffness – were analysed. Although aortic diameter was similar to aortae from WT mice (Fig. 6 A), relative distension (Fig. 6 B), compliance (Fig. 6 C) and E _P (Fig. 6 D) were significantly different in eNOS^−/− mice.

Figure 6. Geometrical and biomechanical parameters of eNOS^−/− mice aortic segments .

Increased aortic stiffness was confirmed in eNOS^−/− mice (n = 5). Although diastolic diameter (D ₀) was similar to wild‐type (WT) mice (n = 5) (A), relative distension (B), compliance (C) and the Peterson modulus (E _P) (D) were significantly different. Measurements were done at a physiological pressure range (80–120 mmHg) and stretch frequency (10 Hz). Line and error bars represent median and interquartile range, respectively. ^** P < 0.01, Mann–Whitney U test.

Functional and physical integrity of the vessel segment is maintained over time

The effect of repetitive stretching on elastin fibre integrity was evaluated on sections of segments after being mounted for 4 h in the ROTSAC set‐up with or without stretching (10 Hz, physiological pressure). No differences in general morphology were observed on H&E stained sections (data not shown). Orcein‐stained aortic sections revealed no difference in elastin fibre integrity after stretching (Fig. 7 A and B) and E _P did not change significantly after a 4 h stretch protocol on aortic segments from both WT and eNOS^−/− mice. These data indicate that repetitive stretching does not damage the elastin fibres in the organ bath.

Figure 7. Effect of repetitive stretch on elastin fibre integrity and E_P .

Orcein‐stained section of aortic segments of WT mice after being mounted in the ROTSAC set‐up at physiological pressure. The integrity of the elastin fibres after 4 h of stretching at 10 Hz (A) was similar to when no stretch was applied (B). E _P at 80–120 mmHg was unchanged (P = 0.97, paired t‐test) (C) after 4 h of stretching in a group of randomly selected mice with different baseline values of E _P (WT, n = 3; eNOS^−/−, n = 5; P = 0.97, paired t test). Scale bar: 100 μm. [Colour figure can be viewed at wileyonlinelibrary.com]

Discussion

In the present paper, we describe the validation of a novel organ bath set‐up that allows the assessment of isobaric stiffness parameters while cyclic stretch is applied at a physiological frequency. This is potentially important as several studies suggest that aortic mechanics are regulated by cyclic stretch in an amplitude‐ and frequency‐dependent matter (Lantelme, 2002; Peng et al. 2003; Tan et al. 2012).

All vertebrates and invertebrates with a closed circulatory system have arteries that show non‐linear elasticity. In the present study, this intrinsic property of the aorta was used to validate the sensitivity of the set‐up to detect changes due to increased distension pressure. Measurements were done at normal (80–120 mmHg) and increased (100–140 mmHg) pressure, while keeping the stretch frequency and amplitude the same. In basal conditions, E _P increased by 20% with increased pressure and the effect of pressure remained significant for all conditions.

Traditionally, passive vessel wall components such as elastin and collagen fibres are considered the most important determinants of arterial stiffness. However, active components that modulate VSMC tone are also known to play an important role (Wilkinson & McEniery, 2004). Therefore, we pharmacologically modulated VSMC tone to test the sensitivity of our set‐up to pick up these changes. Contraction of VSMCs using PE significantly increased isobaric stiffness. When the production of NO was blocked in PE‐stimulated segments, isobaric stiffness increased by 80% as compared to baseline. This large effect of eNOS inhibition can be expected since we reported earlier that, in contrast to muscular arteries, the elastic arteries such as the aorta, produce large amounts of NO in basal conditions (Leloup et al. 2015). This was also suggested by in vivo experiments in rodent models and humans that demonstrated the direct link between (basal) NO production and arterial stiffness (Kinlay et al. 2001; Wilkinson et al. 2002; Schmitt et al. 2005; Isabelle et al. 2012).

To further confirm the ability of our set‐up to pick up differences in aortic stiffness, we used two genetic mouse models. We were able to demonstrate an increased aortic diameter and increased stiffness in segments of the ApoE^−/−Fbn1^C1039G/+ aorta. Indeed, isometric wire myograph analysis demonstrated a steeper stress–strain relationship in 20‐week‐old ApoE^−/−Fbn1^C1039G/+ mice (Van Herck et al. 2009), and histological analysis revealed a lower elastin content, and increased elastin fragmentation and vessel diameter (Van der Donckt et al. 2015).

In aortae from eNOS^−/− mice, stiffness parameters were significantly increased while aortic diameter was similar to aortic segments from WT mice. We and others previously showed increased pulse wave velocity – an important in vivo parameter of arterial stiffness – in eNOS^−/− mice (Soucy et al. 2006; Leloup et al. 2014) and the stress–strain relationship of aortic segments of eNOS^−/− mice was steeper as compared to WT mice (Jung et al. 2013).

These data illustrate the applicability and validity of the ROTSAC set‐up to measure biomechanical properties of isolated mouse aortic segments. In contrast to in vivo analyses of pressure diameter relationships (Giannattasio et al. 2008; Vayssettes‐Courchay et al. 2011) that provide local biomechanical and visco‐elastic information in a physiological environment, the ex vivo ROTSAC set‐up, in which segments of the aorta are isolated, allows us to independently alter the distension pressure, pulse pressure, stretch frequency and composition of the extracellular fluid. Thereby, the effects of VSMC tone, pulse pressure, stretch frequency and distension pressure can be separately investigated to study the active regulation of aortic compliance.

Nevertheless, some disadvantages are inevitably linked to the lack of perfusion. Firstly, similar organ bath set‐ups based on perfusion allow direct measurement of luminal pressure, while this set‐up relies on indirect calculations using the Laplace relationship, assuming a thin, isotropic and homologous wall. The mouse aorta does not meet these assumptions so one should be cautious when interpreting absolute values of calculated pressure. However, the Laplace relationship has been used extensively before in ex vivo biomechanical analysis of the mouse aorta (Syyong et al. 2009; Van Herck et al. 2009; Butlin et al. 2015) and here we confirm that it provides a sufficiently sensitive and practical method to estimate transluminal pressure to evaluate the effects of distension pressure and VSMC tone. Secondly, perfusion‐based set‐ups expose the endothelial cells to shear stress, a well‐known stimulator of endothelial NO production. As mentioned before, we reported that elastic arteries of the mouse produce large amounts of basal NO in vitro. Only very recently we observed that repetitive stretching of the isolated mouse aorta increased basal NO production even more in an amplitude‐dependent manner (unpublished data). Although it is assumed that shear stress is homogeneously distributed in the arterial tree, a review of the literature of shear stress measurements in healthy volunteers revealed that shear stress is lower in the aorta versus other parts of the arterial tree (Pantos et al. 2007). These observations indicate that the contribution of shear stress is more important in the muscular arteries to regulate vessel diameter and, hence, flow, while the elastic arteries – where cyclic deformation is larger – rely more on stretch‐dependent and basal NO production to keep VSMC tone low and the pulse smoothening capacity of the aorta high (Leloup et al. 2015). To determine the relative contribution of basal NO production, cyclic stretch and shear stress‐dependent NO production in the different vessel types, further research is required. Thirdly, upon isolating the aorta, neurogenic and hormonal stimuli are lost and these have been shown to play a role in the regulation of large artery mechanics (Pagani et al. 1975; Barra et al. 1993). A study that compared the effects of vasoactive substances on the aorta found large differences in vivo versus ex vivo (Butlin et al. 2015) and visco‐elastic properties of the aorta change significantly after isolation of the aorta from the animal (Boutouyrie et al. 1997). Therefore, one should be cautious when translating observations from ex vivo set‐ups such as the ROTSAC to the physiological in vivo setting as the combined action of mechanical, neurogenic and hormonal stimuli may differ.

Overall, we were able to show that this set‐up allows fast and straightforward assessment of the intrinsic mechanical and geometrical properties of aortic segments ex vivo. We confirmed the dependency of aortic stiffness on distension pressure and demonstrated large effects of increased VSMC tone on arterial stiffness in aortic segments of WT mice. Two mouse models with known increased arterial stiffness were used to further assess the sensitivity of the set‐up and known differences in vessel mechanics were confirmed. In contrast to (pulsatile) perfusion‐based set‐ups, only a limited amount of tissue is required (1–2 mm aortic segments), allowing parallel testing of multiple conditions, thereby drastically reducing the number of animals required. The aforementioned properties of this set‐up are novel and complementary with perfusion‐based ex vivo set‐ups or in vivo pressure–diameter loop analysis. It will facilitate future research how active vessel wall components (endothelial and VSMC crosstalk) and mechanical stimuli such as distension pressure, pulse pressure and frequency affect arterial biomechanics. More knowledge about the role of these factors in arterial ageing, exercise and hypertension is crucial for the development of specific de‐stiffening strategies that are believed to play a crucial role in future cardiovascular health management.

Additional information

Competing interests

The authors declare that they have no competing interests.

Author contributions

A.L., C.V.H. and P.F. conceived and designed the experiments. Acquisition, analysis, and interpretation of data were carried out by A.L., C.V.H., A.K., S.D.M. and P.F. in the laboratories of W.M., G.D.M., D.S., G.D.K. and P.F. The manuscript was drafted and critically revised by A.L., C.V.H., A.K., S.D.M., W.M., G.D.M., D.S., G.D.K. and P.F. All authors have approved the final version of the manuscript and agree to be accountable for all aspects of the work. All persons designated as authors qualify for authorship, and all those who qualify for authorship are listed.

Funding

A.L. is supported by the Fund for Scientific Research (FWO) Flanders, grant number 11W3716N.