The Morphology of Internal Elastic Lamina Corrugations in Arteries Under Physiological Conditions

Abstract

Background:

In elastic and resistance arteries, an elastin-rich membrane, the Internal Elastic Lamina (IEL), separates the tunica intima from the underlying tunica media. The IEL often appears wrinkled or corrugated in histological images. These corrugations are sometimes ascribed to vessel contraction ex vivo, and to fixation artifacts, and therefore regarded as not physiologically relevant. We examine whether the IEL remains corrugated even under physiological conditions.

Methods:

The diameters of carotid arteries of anesthetized pigs were measured by ultrasound. The arteries were then excised, inflated within a conical sleeve, fixed, and imaged by confocal microscopy. The conical sleeve allows fixing each artery across a wide range of diameters which bracket its ultrasound diameter. Thus the study was designed to quantify how corrugations change with diameter for a single artery, and test whether corrugations exist when the fixed artery matches the ultrasound diameter.

Results:

At diameters below the ultrasound diameter (i.e. when the artery was constricted as compared to ultrasound conditions), the IEL corrugations were found to decrease significantly with increasing diameter, but not fully flatten at the ultrasound diameter. The contour length of the IEL was found to be roughly 10% larger than the circumference of the artery measured by ultrasound. The physiological diameter is likely to be even smaller than the ultrasound diameter since ultrasound was conducted with the animal under general anesthesia which leads to vasodilation, suggesting a higher level of corrugation under physiological conditions. For arterial cross sections constricted below the ultrasound diameter, the IEL contour length decreased roughly with the square root of the diameter.

Conclusion:

The primary conclusions of this study are: a) the IEL is corrugated when the artery is constricted and flattens as the artery diameter increases, b) the IEL is corrugated under physiological conditions and has a contour length at least 10% more than the physiological arterial diameter, and c) the IEL despite being relatively stiffer than the surrounding arterial layers, does not behave like an inextensible membrane.

1. Introduction

Elastic arteries are multi-layered structures: the innermost intima, followed by the media, and the outermost adventitia (Hecker and Duncker, 2021; Thiel et al., 2017).The intima consists of a continuous layer of endothelial cells lining the lumen of the artery. The adventitia consists of collagen and elastin lamellae and anchors the artery to the surrounding tissue (Hecker and Duncker, 2021; Thiel et al., 2017). Between these two layers is the media, which consists of multiple layers of elastin and smooth muscle cells oriented circumferentially, and interspersed with collagen fibers. The number of elastin layers varies with the size and function of the artery. However, all arteries have an internal elastic lamina (IEL) that separates the intima from the media (Hecker and Duncker, 2021; Thiel et al., 2017). This layer is composed of roughly 79% elastin (Schwartz et al., 1981) and is much thicker in large arteries than in arterioles (Schwartz et al., 1981; Thiel et al., 2017). The IEL includes fenestrations of around 3-4 μm in diameter (Sandow et al., 2009) which are needed for mass transport between the intimal and medial layers (Kirby et al., 2013; Sandow et al., 2009).

In most histological cross-sections of arteries, the IEL appears to be heavily corrugated with wavelengths on the order of a few tens of microns (Aiello et al., 2003; Bannister et al., 1975; Greensmith and Duling, 1984; Lee and Chien, 1979; Mayberg et al., 1978; Mello et al., 2004; Movat and Fernando, 1963; Nguyen et al., 2020; Orsi et al., 2006; Orsi et al., 2015; Schwartz et al., 1980; Svendsen and Tindall, 1983; Svendsen and Tindall, 1988; Svendson, 1985; Swinehart et al., 1976; Tindall and Svendsen, 1982). Tindall and Svendsen (Svendsen and Tindall, 1988) have summarized the optical and electron microscopy studies from the 1930’s to 1980’s that revealed the presence of arterial luminal topography. These corrugations are often thought to be an artifact of the fixation process, attributable to the vasoconstriction that occurs when arteries are extracted and preserved (Burton et al., 2017; Choy et al., 2005; Paavilainen et al., 2010). Our study was initiated to determine if these IEL corrugations exist in vivo under physiological conditions (Svendsen and Tindall, 1988) or if they are indeed a fixation artifact.

The potential roles of a corrugated arterial structure are several (Greensmith and Duling, 1984; Svendsen and Tindall, 1988). First, corrugations may be an adaptive mechanism to support the propagation of pulsatile blood flow through the arterial tree by allowing the vessels to cyclically expand and contract. Second, the corrugations may offer an increased surface area for diffusive exchange. Third, corrugations increase the sensitivity of the cross-sectional area to vasoconstriction (Greensmith and Duling, 1984). Thus, due to corrugations, even a modest muscular contraction of the wall can lead to near-complete closure due to “crumpling of wall elements”, especially for small diameters. Finally, our own research (Nath et al., 2020; Nguyen et al., 2020) on synthetic surfaces in contact with blood suggests that surfaces that transition repeatedly from smooth to corrugated (called dynamic topography in our previous research) can avoid surface fouling. A possible mechanism may be delamination driven by the cyclical change in surface topography (Pocivavsek et al., 2018). Thus, the corrugations may serve as a mechanical mechanism of antithrombosis which complements the well-defined antithrombotic properties of the endothelial surface.

This question of whether the corrugations exist in vivo is difficult to address by experiments on live animals. Corrugations on the order of a few tens of microns are too small to detect using current noninvasive imaging techniques such as ultrasound imaging. Invasive techniques such as optical coherence tomography or intravascular ultrasound also lack the necessary resolution. While microCT imaging can reach these resolutions, the dynamic nature of arterial pulsations makes the technique unsuitable for in vivo imaging. Ex vivo studies of fixed arteries allow the corrugated structure to be visualized with exquisite precision, but most fixation processes can induce significant distortion of the tissues. Moreover, arteries undergo vasoconstriction in the setting of hypotension, dehydration, tonus of the arterial wall, hypoxia, hypothermia, or from handling during tissue isolation and fixation (Burton et al., 2017; Choy et al., 2005; Davies and Bowyer, 1975; Huang et al., 2009; McGarvey et al., 1980; Paavilainen et al., 2010; Steinman et al., 2017; Swinehart et al., 1976; Tindall and Svendsen, 1984). Thus, the relationship between the in vivo structure vs the structure observed in ex vivo studies is not clear.

In vitro studies have also examined how the amplitude of the corrugations become smaller as arteries are distended (Aiello et al., 2003; Bannister et al., 1975; Greensmith and Duling, 1984; Lee and Chien, 1979; Movat and Fernando, 1963; Nguyen et al., 2020; Schwartz et al., 1980; Svendsen and Tindall, 1983; Svendsen and Tindall, 1988; Svendson, 1985; Swinehart et al., 1976; Tindall and Svendsen, 1982), and they agree that for arteries ranging from a few mm to less than 100 μm in diameter, the corrugation amplitude increases with decreasing diameter or luminal pressure. This raises the question: What is the situation in vivo? Are the arteries sufficiently distended that they are smooth, or are they insufficiently distended such that they retain significant corrugations under physiologic conditions? In fact, Svendsen and Tindall (Svendsen and Tindall, 1988) specifically discussed this question and summarized the "for" and "against" evidence, but did not come to a definitive conclusion of whether arteries are corrugated in vivo.

Before proceeding, it is useful to clarify the mechanics underlying a corrugated topography. Composed primarily of elastin, the IEL is much stiffer than the underlying medial smooth muscle cell layer. Thus, the near-lumen mechanics is expected to follow that of layered composites comprising a thin stiff layer bonded to a thicker but softer layer (Bowden, 1998). Under compressive loading, the thin stiff layer responds by bending deformations (rather than in-plane compression) and these corrugations grow in amplitude. Indeed, similar stiff-on-soft layer mechanics is also responsible for the wrinkling of skin (Chen et al., 2021; Kruglikov and Scherer, 2018). In the extreme case when the modulus mismatch is large, the stiff layer may behave as an “inextensible” layer, i.e. accommodate the compression entirely by bending without bearing any compressive strain (Pocivavsek et al., 2008). In this limiting case, the topographic corrugations become a geometric effect, i.e. the corrugations solely depend on the diameter of the blood vessel. Accordingly, any mechanism that causes arterial shrinkage such as loss of pressure upon excision or vasoconstriction, would also induce the IEL to become corrugated. That the IEL mechanics follow stiff-on-soft layer mechanics is implicit in the previous literature, although it has generally not been explicitly stated. Indeed, the mechanics of a stiff film on a soft substrate is well-recognized as a common driver of morphogenesis in biology (Ambrosi et al., 2019; Cerda, 2005; Shyer et al., 2017; Shyer et al., 2013).

This physical picture also suggests that the question of physiological relevance may be addressed by quantifying the ratio:

$$\frac{contourlengthofIEL}{physiologicalluminalcircumference}$$

The denominator is the reference value against which the IEL contour length is normalized. A ratio exceeding 1 would imply that the IEL is unlikely to remain smooth under physiological conditions. In fact, a similar ratio was used for quantification by Greensmith and Duling (Greensmith and Duling, 1984) who used the zero-pressure circumference, and by Svendsen and Tindall (Svendsen and Tindall, 1988), who used a luminal circumference calculated from external measurement of in vivo diameter for normalizing (rather than the physiological inner diameter).

Accordingly, our approach is as follows: We measure the physiological diameter in a live animal under anesthesia using ultrasound measurements. We then extract and fix the artery at various diameters. Quantitative image analysis is done to measure the IEL length and compare it against the physiological diameter. To our knowledge, this is the first study on arterial topography that compares ex vivo results against physiological diameter measured in the live animal and hence can judge whether the wrinkles flatten out at a diameter smaller than that in a live animal. One further significant aspect of this research is that we greatly reduce the uncertainty in our measurements by fixing a single artery at a variety of diameters. In contrast, in previous research on this topic, a different artery had to be used at each different diameter. Thus, the reference value used for normalization varied from specimen to specimen, thus leading to greater uncertainty in the measurements.

2. Methods

2.1. Data Collection

Experiments were conducted on the carotid arteries (Left Carotid (LC) and Right Carotid (RC)) of pigs between the weights of 40 and 55 kgs. The weight and sex of each animal is listed in Supplementary Table S1. Pigs were selected for this surgery since these animals were part of other experimental research unrelated to this article which involved surgery, followed by euthanasia. The parts of carotid arteries used for this study were untouched and their health unaffected during the surgery. Accordingly, our research on IEL topography did not raise additional ethical concerns. Further, the relatively large size of the arteries facilitated quantitative image analysis. Information on the surgical procedure and drugs administered to the animals is provided in Supplementary Information Section S.3. The carotid arteries were not the target of that research and were excised immediately after euthanasia. Accordingly, no further ethical approval was needed to conduct these studies.

Artery diameters were measured in vivo using GE Vscan Air^™ ultrasound device. Ultrasound imaging was conducted under general anesthesia (Isoflurane, 2% with 2.5 l/min O₂) and prior to any surgical process. 3-5 ultrasound images at different positions of the carotid artery of each pig were averaged and the corresponding ultrasound-measured diameter will be denoted by $D_{US}$ henceforth. The error in the measurement of $D_{US}$ due to resolution and averaging was estimated to be ± 4%. The various values of $D_{US}$ obtained from each animal are shown in Supplementary Table S1. Some arteries or cross sections were not of sufficient quality for quantitative image analysis and hence a total of 37 cross sections from 9 arteries were used in the analysis.

Immediately after excising from the animal, the arteries were rinsed in saline and the surrounding tissue was surgically removed. The arteries were mounted on the luer connectors in the inflation setup and secured in place. Fig 1(a) is a diagram depicting the inflation setup while Fig 1(b) shows a photograph of the actual setup. 3D printed rigid sleeves with conical inner surface were placed over the artery before inflating (Fig 1(c)). These conical sleeves were designed to have inner diameter ranging from 3 mm on one end to 7 mm on the other. These sleeves ensured that inflation of the covered section of the artery is restricted to the inner diameter of the sleeve. Multiple conical sleeves were sometimes placed upon an artery if the length of the excised artery was sufficiently long (e.g. Fig 1(c) shows an artery with 3 conical sleeves). This allowed us to obtain additional conical sections on the same artery which offered additional spare cross-sections, although only one conical section was sufficient for analysis. One end of the artery was connected to the air outlet while the other end was closed off using a luer stop. The arteries were then subjected to pulsatile inflation cycles for around 1 minute, where the pressure inside oscillated between 0 and 2 psi (approximately 100 mm Hg). The purpose of repeated inflation and deflation of the arteries was to relax the constricted artery to allow distension when pressurized later. The tissue fixation process started within approximately 15-20 mins of excision during which it was kept hydrated. Standard tissue fixation protocol was followed for fixing the arteries under inflation (Choy et al., 2005; McGarvey et al., 1980; Steinman et al., 2017; Swinehart et al., 1976). The arteries were filled with 4% paraformaldehyde, while being immersed in the same solution and inflated to a high pressure of 4 psi for 1 hour. This pressure was regulated so that the arteries had maximum contact with the conical sleeve without damaging the tissue. After 1 hour, the arteries were emptied of the solution, and filled with and immersed in 30% sucrose in phosphate buffer solution (PBS) for 24 hours. The arteries were then taken off the setup and stored at −80° C. Fig 2 shows artery (P10322 LC) fixed using the above method, showing conical segments where the sleeve restricted expansion. The central benefit of this approach is that it allows the same artery to be examined across a range of diameters that encompass the diameter measured by ultrasound on the anesthetized animal. A secondary benefit is that since the arteries are pressed against a circular surface during fixation, their cross-sections remain nearly circular which facilitates further data analysis.

Fig 1:

(a) Conceptual diagram of the inflation setup, (b) Actual experimental setup, and (c) Artery inflated with fixation fluid passing through multiple 3D printed sleeves with conical inner surface

Fig 2:

Fixed arteries (P10322 Right and Left Carotid) showing conical segments where the sleeves were present (indicated by red trapeziums)

10 μm thick sections were taken from the conical part of the frozen arteries. 2-3 sections were taken at each diameter for 4-5 different diameters along the length of the conical part. To observe the effect of vessel diameter on the luminal corrugations, sections of different diameters were imaged. The range of diameters was chosen to bracket the diameter $D_{US}$ observed by ultrasound, and at least one section had a diameter close to $D_{US}$. These included diameters small enough to observe heavy corrugations, to larger ones where the IEL had flattened out. The artery sections were imaged under Nikon A1 spectral confocal microscope at a resolution of ~1.6 μm/pixel. Auto-fluorescence of elastin was used to image the IEL. Scalebars from the images were also converted to XY data for converting pixel data to mm scale. The thickest innermost layer of elastin in the images was considered for IEL topography analysis. Fig 3 shows representative confocal images of artery sections (P723 LC) at different diameters, and the magnified view below each image shows the degree to which the IEL corrugations flatten as the artery diameter increases. We emphasize that all three images were obtained from the same artery.

Fig 3:

(a to c) Confocal images of artery (P723 LC (cross-sections # 1, 3 and 5)) at increasing diameters and zoomed in view (red box) of the IEL

2.2. Data Analysis and Results

Two values were quantified from each image: the degree of constriction with respect to the diameter measured by ultrasound, and the degree of IEL corrugations (the quantity $C_f$ defined below). To do so, the images were first converted to binary and the IEL layer was isolated (Fig 4(a)). Boundaries of the isolated IEL were then converted to XY data. The center of the arterial cross-section was calculated and the XY data was then converted to polar coordinates $(r,\theta )$. The inner and outer boundaries were thus separated and an average radius $(r_{avg})$ was calculated from these data. Fig 4(b) shows the XY data of the isolated IEL along with the circle of radius, $r=r_{avg}$. Fig 4(c) shows the inner and outer boundaries of the corrugated IEL along with $r_{avg}$ when converted to polar coordinates. The average diameter $(D_{avg}=2r_{avg})$ of the arterial cross-section and the non-corrugated circumference $(L_{avg}=\pi D_{avg})$ was then calculated. The various values of $D_{avg},D_{US}$ and their ratio for all the cross-sections are shown in Supplementary Table S1. The supplementary Fig S1 shows the same information of the range of $D_{avg}∕D_{US}$ graphically.

Fig 4:

(a) Isolated IEL binary image, (b) IEL mapped with fitted circle of radius, $r=r_{avg}$ and zoomed in view (red box), and (c) IEL inner and outer boundary mapped in polar coordinates with $r_{avg}$, for a section of the artery (P723 LC, cross-section #1 $(C_f=2.68)$

Contour length of the IEL $(L_c)$ was calculated using the inner boundary of the isolated IEL in polar coordinates (Fig 4(c)). These quantities are shown schematically in Fig 5.

Fig 5:

Schematic of arterial cross section showing calculated quantities $D_{avg}$, $L_{avg}$ and $L_c$

$$L_c=\sum {(\Delta \theta .r)}^2+\Delta r^2$$ (1)

The degree of corrugation then was calculated using the quantity Corrugation factor, $C_f$, defined as the ratio of the IEL contour length to the non-corrugated circumference of the cross-section.

$$C_f=\frac{L_c}{L_{avg}}=\frac{L_c}{\pi D_{avg}}$$ (2)

$C_f>1$ indicated the presence of a corrugated IEL whereas $C_f=1$ indicated that the corrugations had flattened out. The values of $C_f$ for the various cross sections are listed in the last column of Table 1.

Fig 6(a) shows the dependence of the corrugation factor $(C_f)$ for all cross-sections of successfully imaged arteries on the average diameter of the cross-section, normalized by the diameter observed in ultrasound imaging $(D_{US})$. The following section discusses this figure in detail.

Fig 6:

(a) Showing $C_{f}vs{D}_{avg}∕D_{US}$ for all arterial cross-sections (colored dots), (b) A subset of the data magnifying the region with $C_f<1.6$. The dotted line corresponds to an IEL whose contour length remains fixed during compressive deformations.

An alternative analysis of the images was also conducted where the profile of the IEL (Fig 4(c)) was converted into an amplitude profile. From this, the root mean square (RMS) amplitude can be quantified readily. The decrease of the RMS amplitude with increasing normalized diameter $D_{avg}∕D_{US}$ is shown in supplementary Fig S4. That analysis offers an interpretation similar to that from Fig 6.

3. Discussion

As expected from past research(Aiello et al., 2003; Greensmith and Duling, 1984; Lee and Chien, 1979; Nguyen et al., 2020; Svendsen and Tindall, 1983; Svendsen and Tindall, 1988; Tindall and Svendsen, 1982), $C_f$ reduces, and hence the degree of corrugations reduces, as arteries expand (Fig 6(a)). We now address the central question raised in the Introduction of whether arteries are corrugated under physiological conditions. Some of the arterial cross sections are severely compressed with respect to the ultrasound-measured diameter, with average diameter being less than half of $D_{US}$ and $C_f$ values reaching as high as 3. These are of questionable practical relevance, and certainly less relevant to the question of whether the IEL is corrugated under physiological conditions. Accordingly, we now focus only on relatively modest $C_f$ values (Fig 6(b)).

The thin vertical line in Fig 6(b), drawn at $D_{avg}∕D_{US}=1$, represents an arterial cross section whose diameter in the ex vivo fixed state matches the diameter measured by the ultrasound. The data are seen to cross this line with an intercept of approximately 1.1 suggesting that the IEL is 10% longer than the circumference at the diameter measured by ultrasound.

Before discussing the data further, it is important to emphasize that the $C_f$ defined in Eq. 2 is not a direct measure of amplitude, but a measure of how the contour length of the IEL compares with the circumference of the artery as measured by ultrasound. As mentioned in the Introduction, the simplest mechanical picture of the corrugation mechanics is to treat the IEL as inextensible layer that accommodates constriction by bending deformation rather than in-plane compression. In that case, Eq. 2 can be rewritten as

$$C_f=\frac{L_c}{\pi D_{US}}{\left(\frac{D_{avg}}{D_{US}}\right)}^{-1}$$ (3)

Motivated by this, we sought to compare Fig 6(b) against

$$C_f=a{\left(\frac{D_{avg}}{D_{US}}\right)}^{-b}$$ (4)

The dashed blue line corresponds to $a=1$ (or equivalently $L_c=\pi D_{US}$) and $b=1$. This represents an inextensible IEL in which the corrugations disappear exactly at the diameter of the ultrasound experiment. The experimental values clearly deviate upwards from the data in the vicinity of $D_{avg}∕D_{US}=1$ (i.e. the arteries are significantly corrugated as the diameter approaches the ultrasound value), but $C_f$ also has a weaker dependence on $D_{avg}∕D_{US}$ than expected. By trial and error, we find that $a=1.1$ and $b=0.5$ give reasonable fits to the data. Other pairs of parameters, e.g. ($a=1.06$; $b=0.6$) and ($a=1.12$; $b=0.4$) are shown as thin gray lines in Fig 6(b); parameters outside of these bounds no longer represent the data even qualitatively. Values of $a$ less than 1.06 fall below all the rightmost points in Fig 6(b) regardless of the value of $b$ suggesting that any reasonable fit must have $C_f>1$ as $D_{avg}$ approaches $D_{US}$. While one may consider other variations to Eq. 3 which allow more than two parameters, these do not capture the data better than Eq. 4.

The fact that values of $a$ less than 1.06 are unable to capture the data suggests that the IEL is longer than the circumference corresponding to the ultrasound-measured diameter. Further, only values of $b$ that are significantly less than 1 can capture the data, suggesting that the IEL is not strictly inextensible. This is seen more directly in the Supplementary Fig S2. which directly plots the measured $L_c$ vs diameter for each of the arteries examined and shows that $L_c$ increases with increase in cross section diameter.

While similar data have been published by previous researchers, none sought to reference the measurements of topography to the diameter measured in vivo. For example, Lee and Chien (Lee and Chien, 1979) held canine carotid arteries at various pressures during fixation. The lumen was heavily corrugated at 0 mm Hg and these corrugations decreased in amplitude with increasing luminal pressure (and hence increasing diameter). Similarly, our own group (Nguyen et al., 2020) reported measurements of pig carotid arteries. As in Lee et al, the arteries were fixed while being held at various luminal pressures. The results were qualitatively similar to Lee et al. Finally Greensmith and Duling (Greensmith and Duling, 1984) examined arterioles, but unlike Lee and Chien (Lee and Chien, 1979) and Nguyen (Nguyen et al., 2020), all arteries were fixed at the same pressure. However, the vessel constriction was tuned not by varying luminal pressures but by using the vasoconstrictor (Greensmith and Duling, 1984). Corrugations appeared when vessels were 110-125% of their unstressed diameter, and the arteries became highly corrugated when they were less than 90% of their unstressed diameter. In these articles, the lack of confidence in the in vivo diameter means that although the existence of corrugated arterial topography was well-quantified, it was not possible to comment on the situation in vivo. Svendson and Tindall (Svendsen and Tindall, 1983) quantified the corrugations of the rabbit aorta by measuring the contour length of the IEL in fixed arterial cross sections. They compared the contour length to the in vivo luminal circumference estimated from the in vivo outer diameter of the exposed aorta in anesthetized rabbits. The IEL circumferential length was reported to be 10-30% more than the approximated in vivo luminal circumference. In this paper, we report the IEL to be roughly 10% longer than the measured in vivo circumference, or equivalently, the diameter needed for flattening the IEL exceeds $D_{US}$ by about 10%. Some of this difference between our results vs Svendsen and Tindall may be due to using a different reference configuration for comparison: we have used the ultrasound measurements in vivo diameter whereas Svendsen and Tindall used unstressed in-vitro diameter of excised blood vessel. Of course, some of the difference may also be intrinsic to the difference in the vessel used (rabbit aorta vs pig carotid).

Our observations imply that the luminal surface has modest corrugations when the artery at the ultrasound diameter. To determine whether the corrugations are significant at physiological diameters, we must judge whether the diameter $D_{US}$ observed in ultrasound truly represents the physiological in vivo diameter. It has been well demonstrated that vasodilation occurs during anesthesia (Akata, 2007; Iida et al., 1998; Maud et al., 2018; Ryu et al., 2020; Slupe and Kirsch, 2018; Sullender et al., 2022). Thus, the arterial diameters measured by ultrasound in our study are likely to be larger than the physiological diameter in non-anesthetized animals. Smaller arterial diameter corresponds to a larger $C_f$ and hence more prominent corrugations in the IEL topography. Hence, at physiological diameters, the IEL would be more corrugated than that at the larger diameters we observed in this study.

As mentioned in the Introduction, Svendsen and Tindal have previously summarized the "for" and "against" arguments on the question of whether arteries are corrugated under physiological conditions. Since we have directly measured the in vivo diameter, our observations provide strong evidence on the "for" side of that question. To unambiguously settle this question, it would be essential to directly image the inside of the artery; this may become possible in the future as OCT resolution improves (De Boer, 2009; Fujimoto et al., 2000; Gupta et al., 2022; Ong et al., 2022).

4. Conclusion

In summary, this study examines luminal topography of pig carotid arteries as judged by the corrugated appearance of the internal elastic lamina (IEL). Compared to previous research in this area, our study was designed with two noteworthy aspects. First, each excised artery was fixed while being held inflated inside a conical sleeve, and hence cross sections of each artery could be imaged across a wide range of diameters. Second, the diameter of each artery was measured by ultrasound prior to excision while the animal was anesthetized. Taken together, these two aspects offer confidence on how the luminal topography of a single artery varies with constriction, and what was the diameter of that same artery in the live animal. We draw three central conclusions. First, consistent with past research in this area, the arterial corrugations flatten out as arterial diameter increases. Second, the arterial cross sections corresponding to the ultrasound-measured diameter have an IEL whose circumferential contour length is roughly 10% larger than the circumference calculated from the ultrasound diameter. Since the arterial diameter in the normally-functioning animal is likely to be smaller than the ultrasound diameter measured under anesthesia (since anesthesia induces vasodilation), our results suggest that these carotid arteries are corrugated under physiological conditions. Finally, the IEL does not appear to behave as an inextensible membrane, instead, its circumferential contour length reduces with constriction. Broadly, these results suggest that arteries should not be regarded as having a smooth lumen under physiological conditions.