© 1999 by European Society of Cardiology
Copyright © 1999, European Society of Cardiology
The influence of aging and aortic stiffness on permanent dilation and breaking stress of the thoracic descending aorta
aDepartment of Cardiology, Academic Medical Center, Amsterdam, The Netherlands
bDepartment of Medical Physics, Academic Medical Center, Amsterdam, The Netherlands
cDepartment of Cardiovascular Pathology, Academic Medical Center, Amsterdam, The Netherlands
dDepartment of Cardiology, Leiden University Medical Center, Leiden, The Netherlands
eInteruniversity Cardiology Institute of the Netherlands, Leiden University Medical Center, Leiden, The Netherlands
fBio Implant Services, Leiden University Medical Center, Leiden, The Netherlands
* Corresponding author. Tel.: +31-20-566-2193; fax: +31-20-566-6809 m.groenink{at}amc.uva.nl
Received 10 July 1998; accepted 2 February 1999
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Abstract |
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 Top Abstract 1 Introduction2 Methods3 Results4 Discussion5 ConclusionsReferences |
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Objective: To assess the influence of aging and aortic stiffness on the extent of irreversible deformation and breaking stress of the human thoracic aorta. Methods: From 14 human heart valve donors without aortic disease (mean age 35 years, range 8–59 years), 14 intact segments of the thoracic descending aorta were studied within 48 h after cardiac arrest. In an experimental setup, the segments were submitted to increasing hydrostatic pressure loads, both statically and dynamically, while radius and wall thickness were monitored echocardiographically. Pressure–radius curves were constructed. Radius and wall thickness were determined at a pressure of 100 mmHg. Radius at elastin resting length and collagen recruitment pressure (Pcol, mmHg) were derived from the pressure–radius relationship and stress–strain curves were constructed to yield Young’s moduli of elastin and collagen. Distensibility (D, mmHg–1) was determined while loading the segment with a sinusoidal pressure wave of 120/50 mmHg at both 0.5 and 1 Hz. Subsequently increasing static pressure loads of 400, 800, 1200 and 1600 mmHg were applied. After each pressure load, the increase in aortic radius at a pressure of 100 mmHg (Rinc) was determined. The experiment continued until rupture occurred and breaking stress (
break, N m–2) was calculated. donor age and aortic stiffness were correlated with Rinc and break of the aortic segments. Results: Mean breaking stress of the 14 segments was 2.7x106 N m–2. Breaking stress was negatively correlated with age (r2=0.66) and positively with D (r2=0.44) and with Pcol (r2=0.18). Seven segments survived a pressure load of 800 mmHg, in these vessels, the extent of irreversible dilation was positively correlated with age (r2=0.42) and negatively with D (r2=0.40) and Pcol (r2=0.40). Conclusion: Permanent deformation and rupture of the human thoracic aorta following pressure overload are influenced by age, distensibility and collagen recruitment pressure.
KEYWORDS Arteries; Atherosclerosis; Blood pressure; Remodelling; Ultrasound
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1 Introduction |
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TopAbstract1 Introduction2 Methods3 Results4 Discussion5 ConclusionsReferences |
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Fatal dissection and rupture of the thoracic aorta account for approximately one percent of all deaths [1]. Besides connective tissue diseases and hypertension, aging plays a major role in the process of aortic degenerative disease [1–19]. Although the mechanism is still not fully clarified, ‘senile’ aortic dilation has been attributed to a combination of repetitive cyclic stress [20] and atherosclerotic degeneration [14,19]. In a theoretical model, ‘wear and tear’ of the aortic wall during aging may result in fragmentation of the medial elastin network and repair with collagen, resulting in a change in relative concentrations of fibrillar material [14,15,18]. Injury to medial fibers may be followed by mononuclear cell infiltration and matrix metalloproteinase-mediated elastolytic and collagenolytic activity, resulting in further breakdown of medial fibers [9,10,21–23]. Breakdown of elastic fibers may lead to aortic dilation, whereas collagenolytic activity may cause increased fragility of the aorta [21,24]. Both elastolytic activity [22] and expression for metalloproteinases and elastase in aortic smooth muscle cells [23] have been shown to increase in normal aging aortas. Breakdown of medial fibers and aortic dilation causes increased aortic stiffness [14,18,19,21,24] resulting in increased pulse pressure and wall stress. Concomitant hypertension may further increase the aortic wall stress, and contribute to the development of aortic aneurysms [7,25]. Aortic dilation, increased aortic stiffness and elevated blood pressure are probably responsible for the high wall stress generated in the fragile aorta in the elderly [13–19,21] resulting in further dilation and medial fibrillar damage until dissection and/or rupture takes place. Although the aorta is considered to become more susceptible to dilation and fragile during aging, this has never been demonstrated. Breaking stress of the aorta in the elderly has been studied in tensile force line-ups with aortic strips, assuming isotropic behaviour of the aortic wall [26,27] but not in intact segments. To our knowledge, the extent of permanent dilation achieved after pressure overload of the aorta has never been studied. Aortic stiffness may reflect the functional state of the normal aging aorta and be of prognostic value for the occurrence of aortic dilation and dissection or rupture. The use of intact aortic segments enables the assessment of both static and dynamic viscoelastic behaviour in relation to breaking stress and susceptibility to dilation.
Accordingly, the aim of the present study was to investigate breaking stress and permanent dilation of intact segments of the normal thoracic aorta in relation to age and aortic stiffness.
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2 Methods |
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TopAbstract1 Introduction2 Methods3 Results4 Discussion5 ConclusionsReferences |
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2.1 Aortic tissue
From 14 human heart-valve donors, without any known aortic disease, the thoracic descending aorta was harvested as part of a ‘non-beating heart’ donor program for transplantation purposes. The donor selection was made according to the European Standards for cryopreserved heart valve homografts and with permission for experimental use of the donor tissue (informed consent). Explantation took place within 24 h of circulatory arrest. During harvesting of the heart, pulmonary arteries with bifurcation and aortic arch, the descending aorta was ligated one centimeter distal to the left subclavian artery and fixed at its in situ length. Next, the descending aorta was carefully cleared from the surrounding tissue. The intercostal arteries were cut at a length of approximately 1 cm. Finally, the in situ length of the aortic segment was measured (Table 1) and the segment was explanted. Blood clots were removed by carefully rinsing the aortic segment in 500 ml of sterile 0.9% NaCl. Fascia and fatty tissues were carefully removed from the aortic wall until a smooth surface was achieved.
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2.2 Apparatus and protocol
The segments were studied within 48 h of circulatory arrest. Two rings of approximately 2 mm were taken from either side of the segment for light microscopic studies. Fourteen segments with an in vivo length of 65±5 (SD) mm remained. The stumps of the intercostal arteries were ligated. The segments were stretched to their original in situ length and fixed in an experimental setup (Fig. 1) by means of glass cannulae, adjustable for small variations in segment length. The cannula at the proximal side of the segment was connected to a perfusion reservoir filled with a 0.9% saline solution, and the cannula at the distal site was closed. The cannulae and the segment were filled with the 0.9% saline solution. Air pressure was directed towards the perfusion reservoir by means of a manual pressure controller (Fairchild, model 65A) either directly, or via an electric pressure controller (Fairchild, model T5200), guided by a sinusoidal generator (Krohn and Hite, model 5700). In this way, both static and dynamic hydrostatic pressure could be exerted on the segments. A pressure catheter, connected to a specially adapted pressure transducer (Baxter, disposable PX-600F), was placed inside the segment. This transducer was modified in order to measure high-pressure levels by attenuation of the output voltage. Furthermore, the security valve against pressure overload was made dysfunctional. Pressure measurements with this transducer were linear and reproducible for pressures up to around 1600 mmHg. After amplification and filtering (Bridge amplifier DPM-612A), the signals were recorded by a thermal pen recorder (WEKAgraph, model WK 250R). A 5-MHz ultrasound probe (Hewlett Packard) was positioned and fixed above the aortic segment, imaging continuously the segment.
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The aortic segments were ‘accommodated’ for 30 min with a sinusoidal pressure oscillation at 0.5 Hz of 120/50 mmHg, thus eliminating the effects of smooth muscle cells on elastic behavior [28]. Then, the experiment started with investigation of viscoelastic properties of the segment with a dynamic pressure oscillation of 120/50 mmHg at 0.5 Hz and at 1.0 Hz. Outer diameter variation was measured by means of M-mode echography.
Next, static pressure was exerted on the segments starting at 20 mmHg and increasing in steps of 20–40 mmHg, until a pressure of 400 mmHg was reached. The pressure was kept at 400 mmHg for 20 min. This procedure was repeated to end-pressures of 800, 1200 and 1600 mmHg, and each time starting at a pressure of 20 mmHg, until rupture of the segment occurred. The outer diameter of the segment was continuously measured using two-dimensional mode (2D) echography. After mechanical testing the wet weight of the segment was determined (Sartorius, model H120).
2.3 Histology
Of each of the 2-mm aortic rings at least two representative full thickness arterial wall blocks were taken. The tissue blocks were fixed in 4% formalin solution, neutrally buffered (pH=6.8–7.2), processed through methylated alcohol (50–100%) and embedded in paraffin wax. Six-µm thickness sections were stained with haematoxylin and eosin and elastic van Gieson stain. All sections were screened for the presence or absence of significant pathologic alterations.
2.4 Calculations
Aortic wall volume was calculated from the weight (g) of the segments by means of the equation:
 | (1) |
From external radii and volume data, the internal wall radius (the distance from the luminal center to the luminal edge of the segment) was calculated by means of the equation:
 | (2) |
Wall thickness (h100, mm) was calculated by subtracting the internal radius from the external radius at a pressure of 100 mmHg.
Midwall radius (mm) was calculated by means of the equation:
 | (3) |
The radius–wall thickness ratio (R/h) was determined at a pressure of 100 mmHg.
Per segment pressure–radius curves were constructed (Fig. 2A).
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Radius at resting length of elastin fibers (R0e, mm) was calculated from the x-intercept of the pressure–radius relationship in the linear low-pressure range.
Collagen recruitment pressure (Pcol, mmHg) was defined as the pressure at which the pressure–radius relationship deviated from its initial linear behavior.
Strain was calculated as the radius increase following each pressure step divided by R0e.
Circumferential wall stress () was calculated by means of Eq. (4) [29]:
 | (4) |
=wall stress (N m–2) and P=pressure (N m–2). Stress–strain curves were constructed to yield (Fig. 2B):
- • Young’s modulus of elastin (Ee, N m–2), calculated as the slope of the first (linear) part of the stress–strain curve, and
- • Young’s modulus of collagen (Ec, N m–2), calculated as the slope of the linear part of the stress strain curve in the range 3500 to 5000 N m–2, relative to Young’s modulus of elastin (Fig 2B). Consequently, Ec was corrected for contribution of elastin fibers to total elastic behaviour in the high-pressure range [29]. Hence, Ec=slope–Ee.
- • Distensibility (D, mmHg–1) was calculated from the dynamic pressure–radius relation by means of the equation:
in which
(5) 
R=midwall radius change, Rmin=smallest midwall radius and P=pressure change (mmHg).
- • Maximum achievable strain (Smax), relative to R0e,was determined.
- • Deformation (Rinc, mm) was assessed from the difference between the radius at 100 mmHg before and after each static pressure loading.
- • Young’s modulus of collagen (Ec, N m–2), calculated as the slope of the linear part of the stress strain curve in the range 3500 to 5000 N m–2, relative to Young’s modulus of elastin (Fig 2B). Consequently, Ec was corrected for contribution of elastin fibers to total elastic behaviour in the high-pressure range [29]. Hence, Ec=slope–Ee.
The wall-stress at which rupture occurred, was called ‘breaking stress’ (break, N m–2).
2.5 Statistics
Linear regression analysis was used to correlate age of the donors with R0e, R/h, Ee, Ec, Pcol, D, Smax, Rinc and break of the aortic segments. Furthermore, Rinc and break were correlated with Ee, Ec, Pcol and D. Correlation between parameters was expressed as the least squares best-fit (r2). Confidence intervals (95%) of the regression lines were calculated using Graphpad Prism® v2.01 for Windows®.
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3 Results |
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TopAbstract1 Introduction2 Methods3 Results4 Discussion5 ConclusionsReferences |
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3.1 Histology
All aortic specimens showed a regular organization of lamellar units in the media. In three segments of older donors, there was minimal fragmentation of elastic fibers, but cystic media necrosis was never observed. The intima of the aortas showed diffuse thickening composed of extracellular matrix, increasing in thickness with age. In five segments, obtained from donors >40 years of age, atherosclerotic plaques, composed of fibrosclerotic tissue and extracellular lipid pools were found. In some of these, there was a slight attenuation of the underlying media. Ulcerating atherosclerotic or inflammatory disease of the aorta was never observed.
3.2 Basic characterization of findings
Distensibility at 1 Hz correlated well with distensibility at 0.5 Hz (Fig. 3A). Distensibility at 1 Hz was subsequently used for further analysis. Collagen recruitment pressure correlated well with distensibility (Fig. 3B). A very moderate permanent dilation to approximately 105% of the initial radius could only be achieved in the seven aortic segments that could withstand a pressure load of 800 mmHg. In nine segments aortic rupture was preceded by aortic dissection. In these segments, rupture occurred immediately after dissection. Rupture of the aortic segments occurred mostly at the ostium of an intercostal artery, perpendicular to the long axis of the segment (Figs. 4 and 5).
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3.3 Effect of aging on aortic dimensions and mechanical properties
Baseline characteristics of the aortic segments and calculated entities are summarized in Table 1. Fig. 6A and 6B show typical examples of the effect of aging on pressure–radius (Fig. 6A) and stress–strain (Fig. 6B) curves. With increasing age, aortic radius at elastin resting length increased and collagen recruitment pressure (black dots in Fig. 6A) decreased. Age correlated positively with radius at elastin resting length (Fig. 6C), y=0.062x+5.2 and negatively with radius–wall thickness ratio (Fig. 6D), y=–0.24x+22, collagen recruitment pressure (Fig. 6E), y=–1.14x+117, and distensibility (Fig. 6F), y=–0.0001x+0.001. Although the radius at elastin resting length increased with age, radius–wall thickness ratio at a pressure of 100 mmHg actually decreased with age. As already suggested by Fig. 6B, no correlations between age and Young’s moduli of either elastin or collagen could be demonstrated (Fig. 6G, H). All aortic segments could be loaded up to a pressure of approximately 250 mmHg, after which only slight further deformation could be achieved (Fig. 7A). Age correlated negatively with maximum achievable strain (Fig. 7B), y=–0.010x+0.78, positively with degree of permanent deformation (Fig. 7C), y=0.02x–0.07, and negatively with breaking stress (Fig. 7D), y=–0.075x+5.3.
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3.4 Aortic stiffness, deformation and rupture
The degree of permanent deformation,following a pressure load of 800 mmHg, correlated positively with distensibility (Fig. 8A), y=162x+1.49, and negatively with collagen recruitment pressure (Fig. 8B), y=–0.015x+1.9. Breaking stress was positively correlated with distensibility (Fig. 8C), y=424x+0.58 and, to a lesser extent, with collagen recruitment pressure (Fig. 8D), y=2.6·103x+6.9·105. No correlations between Young’s moduli of elastin or collagen and either breaking stress or degree of permanent dilation could be demonstrated.
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4 Discussion |
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TopAbstract1 Introduction2 Methods3 Results4 Discussion5 ConclusionsReferences |
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With increasing age, aortic dimensions change and aortic stiffness increases, as has been demonstrated in several previous studies [13,15–18]. In the present study, we also showed an age-related increase in fragility and susceptibility to permanent dilation of the thoracic aorta. Although degenerative changes were only mildly present in the wall of the aortic segments from older donors, as shown histologically in our study, it is concluded that these changes are responsible for alterations in mechanical behaviour.
4.1 Aortic stiffness
There were clear correlations between age and collagen recruitment pressure, and between age and distensibility of the aorta. The importance of age-related shift of the collagen recruitment pressure to lower pressure ranges for total elastic behaviour of the aorta has been shown in previous studies [16–19]. Distensibility correlated well with collagen recruitment pressure and may be of clinical value to estimate the functional state of the aorta in aortic degenerative disease because distensibility can be determined non-invasively [30,31]. No correlations, however, could be shown between age and either Young’s modulus of elastin or Young’s modulus of collagen, raising doubt about the clinical usefulness of these parameters [29].
4.2 Aortic deformation
The older the aortic segments, the less they could be stretched. When stretched beyond a certain degree, permanent dilation occurred. The degree of permanent dilation correlated with age, collagen recruitment pressure and distensibility in seven segments that could withstand a pressure load of 800 mmHg. Although it has been suggested from clinical experience and previous studies [7,25] that permanent aortic dilation may be achieved merely by force, only a very limited permanent dilation was achieved by means of hydrostatic pressure augmentation. In the pressure range above 250 mmHg, where all collagen fibers are recruited, there seems (at least in non-diseased aortic segments) to be an ‘all or nothing’ mechanism, hampering further dilation until dissection and/or rupture occurs. It can not be excluded that a more significant dilation can be achieved by loading aortic segments for several hours, days, or even months.
4.3 Aortic rupture
Mean breaking stress in the present study was (2.7 ±1.5)·106 N m–2, which corresponds reasonably well with results from Raghavan et al. [26] who reported a mean breaking stress of the infrarenal aorta of (2.0±1.0)·106 N m–2. In that study, the authors used aortic strips from seven cadaveric organ donors aged 47±4 years to test breaking stress in a tensile force line-up. Restricting the breaking stresses of segments in the present study to the same age group (six segments), a mean breaking stress of (1.7±0.4)·106 N m–2 was calculated, which is slightly lower than in Raghavan’s study. For a number of reasons, such as composition and elastic properties [32] one would expect breaking stress of the thoracic aorta to exceed the breaking stress of the infrarenal aorta. In our study intact segments were tested, bearing loads in all directions, instead of circular or longitudinal loading of aortic strips, which may partly explain the small discrepancy between these results. In the majority of our cases, rupture was preceded by dissection, mostly occurring at the location of an intercostal artery, which can only be demonstrated in intact segments. Moreover, if the aorta would behave as an isotropic body, one should expect rupture to occur longitudinally (perpendicular to the direction of maximum deformation), which was not shown in our study. In a study by Adham et al. [27] 46 strips from the human thoracic descending aorta were stretched until breaking point. Although we are not informed about the age of the donors in Adham’s study, a mean breaking stress of (2.0±0.6)·106 N m–2 tallies well with our results and those from Raghavan et al. [26] suggesting no difference in breaking stress between the infrarenal and the thoracic descending aorta. However, we showed that breaking stress of the aorta is strongly influenced by age.
Breaking pressures found in the present study seem highly unphysiological. However, wall stresses of the same magnitude as the breaking stresses in the present study can be generated at physiologic pressures in aortic aneurysms, or during isometric tension sports as weight lifting [33]. Breaking stresses in aortic aneurysms are probably even much lower than those in normal aortas [26] and the present study supplies values for wall stresses, which should certainly not be exceeded in a patient with an aortic aneurysm.
We demonstrated correlations between aortic stiffness (collagen recruitment pressure and distensibility) and both susceptibility to permanent dilation and breaking stress of the thoracic aorta. This may be a rationale for further studies concerning the predictive value of in vivo measurable aortic stiffness in patients with aortic disease.
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5 Conclusions |
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TopAbstract1 Introduction2 Methods3 Results4 Discussion5 ConclusionsReferences |
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- 1. The effects of normal aging on the thoracic descending aorta are: (a) a decrease in radius–wall thickness ratio at a pressure of 100 mmHg, (b) a reduction in maximum achievable strain, (c) weakening of the aortic wall, and (d) an alteration in elastic behavior, characterized by a shift to lower pressures of the collagen recruitment pressure and reduced distensibility.
- 2. The holding strength and permanent dilation following pressure overload of the thoracic descending aorta are correlated with age, collagen recruitment pressure and distensibility.
- 3. Forced rupture of the thoracic descending aorta occurs suddenly, without a further change in dimensions beyond a certain wall stress, mostly preceded by dissection originating in the ostium of one of the intercostal arteries.
- 4. Rupture behavior of the thoracic descending aorta seems to be strongly influenced by anisotropy, which may hamper investigation with traction tests on aortic strips.
- 2. The holding strength and permanent dilation following pressure overload of the thoracic descending aorta are correlated with age, collagen recruitment pressure and distensibility.
Time for primary review 14 days.
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Notes |
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1 M. Groenink is supported by a grant from the SORBO Heart Foundation.
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TopAbstract1 Introduction2 Methods3 Results4 Discussion5 ConclusionsReferences |
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