International Congress Series 1274 (2004) 64 – 74

www.ics-elsevier.com

Elasticity imaging of atheroma with transcutaneous
ultrasound both in longitudinal-axis andshort-axis planes
Hiroshi Kanai*, Hideyuki Hasegawa, Nozomi Nakagawa
Graduate School of Engineering, Tohoku University, Aoba-ku Aramaki-aza-Aoba 6-6-05, Sendai 980-8579,
Japan

Abstract. This paper describes a noninvasive method for evaluating regional viscoelasticity of thearterial wall in which a novel method is applied to measure minute changes in wall thickness causedby the change in internal pressure. (1) By comparing the pathological findings with the elasticity
distribution, statistical parameters for lipid and a mixture of smooth muscle and collagen fibers canbe determined. By applying the method to the common carotid arteries, thin collagen fiber was
clarified and soft inclusion of lipid in plaque was detected. (2) A new beam steering method was
proposed to apply the above method to the plane, which is perpendicular to the axis of the artery. (3)For improvement of classification of tissue, a method was proposed to measure viscosity bygenerating the change in internal pressure using the external actuation from the skin surface, namely
remote actuation. These methods offer the potential for detection of plaque vulnerability in a clinicalsetting. D 2004 Elsevier B.V. All rights reserved.
Keywords: Viscoelasticity of arterial wall; Tissue characterization; Ultrasound; Atherosclerosis

1. Introduction
Rupture of atherosclerotic plaque is probably the most important factor underlying thesudden onset of acute coronary syndrome [1]. Direct characterization of the compositionand vulnerability of atherosclerotic plaque, rather than of the angiographic lumen [2],may offer insight into the mechanisms of plaque regression and progression [3,4] andthereby promote evaluation of cholesterol-lowering therapy [5,6] for reduction of* Corresponding author. Tel.: +81 22 217 7078; fax: +81 22 263 9444.
E-mail address: hkanai@ecei.tohoku.ac.jp (H. Kanai).
0531-5131/ D 2004 Elsevier B.V. All rights reserved.
doi:10.1016/j.ics.2004.07.029

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cardiovascular events. Magnetic resonance imaging (MRI) and intravascular ultrasonography (IVUS) are promising technologies for directly imaging plaque morphology[7,8]. For the evaluation of dynamic mechanics, arterial elasticity has been determinedby measuring pulse wave velocity [9] and rough change in the diameter of the artery[10,11]. However, the mechanical property of plaque cannot be directly determined bythese methods because they evaluate average elasticity of the entire circumference. A
method to detect the vulnerability of atherosclerotic plaque with sufficient accuracy has
not yet been reported. The purpose of the present study was to determine the regional
viscoelasticity in the arterial wall using commercially available transcutaneous ultrasound equipment.

2. Methods
2.1. Measurement of change in thickness of the arterial wall
An ultrasonic beam was sequentially scanned at M positions with a linear-typeultrasonic probe of 7 MHz using conventional ultrasound diagnostic equipment (ToshibaSSH-140A), and multiple (N m +1) points were preset from the luminal surface to theadventitia along the m-th ultrasonic beam (m=1, . . ., M) with constant intervals of h 0=375
Am at a time t 0 just before the ejection period. By dividing the arterial wall into multiple
layers, we defined the n-th layer (n=1, . . ., N m ) as being between two contiguous points, n
and n+1, along each beam. For measurement of the change in thickness of each of the N mlayers, the instantaneous depth x m,n (t) of the n-th point along the m-th beam wassimultaneously tracked by applying the phased tracking method [12,13] to the received
ultrasound. The minute decrease of several tenths of a micrometer in thickness of the n-thlayer, resulting from the arrival of the pressure wave at the beginning of the ejection
period, was determined by Dh m,n (t)=x m,n+1(t)x m,n (t)h 0.
In the phased tracking method, for calculation of the autocorrelation function
between the quadrature-demodulated signals of sequentially received echoes, a minutephase change of about 0.48 caused by movement of the n-th point during the pulse
transmission interval DT (=200 As) can be accurately determined by introducing a
constraint, namely that their waveforms are identical but their phase values change[12,13]. The lowest value of the change in thickness was validated as being about 0.5
Am by expanding a rubber plate in a water tank [14]. Such a minute change in thickness
cannot be measured by any other method. This method has already been applied to thein vivo detection of the regional instantaneous displacement and change in thickness,with sufficient reproducibility, in the interventricular septum [12,13,15] and the common
carotid arteries [16,17].
2.2. Elasticity estimation
From the ratio of the maximum decrease in thickness during one heartbeat,
Dh m,n,max=maxt |Dh m,n (t)|, to the initial thickness h 0 of the n-th layer, the maximum
deformation of the n-th layer was obtained by De m,n,max=Dh m,n,max/h 0. Because the

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deformation was sufficiently small and was in the linear regime, it showed incremental
strain in the radial direction. By assuming that the arterial wall is incompressible andthat the blood pressure is applied normally to each layer, the elastic modulus of the n-thlayer along the m-th beam, E hm,n , is approximately given by [18]!
1 qm;n;0
Nm  n þ 1
DpEh;m;n iþðn ¼ 1; . . . ; Nm ; m ¼ 1; . . . ; MÞ;2
Nmh0
Dem;n;maxð1Þ
where q m,n,0 is the initial inner radius of curvature of the n-th layer along the m-th beam
at a time t 0. We assumed that the pressure in the arterial wall decreases linearly with the
distance from the intimal side to the adventitia and that the arterial wall is almost
isotropic [19].
For the region with a length of 18 mm (M=60) along the axis of the artery, theregional elasticity E hm,n was estimated on the cross-sectional image. Because thereflected ultrasound was received at a sampling interval of 100 ns (=75 Am along depth
direction) after the quadrature demodulation, we further divided each layer with athickness of h 0 into five points, shifted the initial depth of each layer by 1/5 of h 0 andapplied the above procedure to each depth. Thus, E h,m,n was estimated at intervals of 75
Am in the depth direction and 300 Am in the axial direction. Using a silicone rubber tube
with two layers set in an artificial circulation system, the accuracy of the measurement ofregional elasticity for each layer has already been validated to be about 0.1 MPa [18], thatis the error is about 8% of the elasticity value obtained by a separate static pressure–diameter test.
In in vivo experiments before, the surgical extraction of an iliac artery, and in in vitro
experiments (described below) after such extraction, the average elasticity was between
0.96F0.48 and 0.89F0.31 MPa, respectively; the difference between them being about 8%.
Thus, the slight influence of assaying the artery through the skin was eliminated.
In separate in vivo experiments, when the pressure of the ultrasonic probe on the skinsurface of a healthy subject was set as 8.0, 14.0, 33.2, 40.7, 54.5, and 74.3 mmHg, themeasured diameter of the same point of the common carotid artery changed as 6.8, 6.8, 6.6,
6.1, 6.0, and 5.4 mm, respectively. For higher pressure, the cross-section of the arterychanges from a circular to an oval shape. We confirmed that the measured elasticity is not
influenced by the pressure on the ultrasonic probe to the skin surface as long as the pressureis equal to or less than 30 mmHg. In our in vivo experiments, the ultrasonic probe was heldon the skin surface at a pressure of 30 mmHg.
2.3. Beam steering for measurement in the short-axis plane
As shown in Fig. 1(a), in the long-axis plane, all ultrasonic beams become
perpendicular to the wall by linear scanning. However, only one beam, which passes
through the center of the artery, is perpendicular to the short-axis plane. For
measurement in the short-axis plane, as shown in Fig. 1(b), the ultrasonic beams aretransmitted to the M (=N+1) directions and designed so that each beam passes throughthe center, O, of the artery. Under such a setting, the ultrasonic beams are always

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Fig. 1. (a) Illustration of long-and short-axis planes. (b) Schematic diagram of beam steering.

perpendicular to the arterial wall if the cross section of the artery can be assumed acircle. The beam scanning is designed as follows [20]: The distance, d k , from thecenter, c k , of the transmit aperture of the k-th beam (k=N/2, . . ., 0, . . ., N/2) to thecenter of the probe is expressed as follows:
dk ¼ 0:4 k ½mm:ð2Þ

The beam angle, h k , of the k-th beam is expressed as follows:
hk ¼ arctandk½rad;
Li

ð3Þ

where L i is the distance from the surface of the probe to the point O. In our system, L i isvariable from 8.5 to 21.5 mm with a pitch of 1 mm (i=1, 2, . . . , 14). The focal distance of
the k-th beam is defined byfk ¼

Liþ 9 ½mm:
coshk

ð4Þ

In Eq. (4), the distance from c k to O is added by 9 mm so that the focal depth becomesdeeper than the posterior wall. The beam width at the posterior wall is widened by thissetting, and a wider beam is robust against wall motion, which slips off the beam [21]. Bychanging the distance, L i , each ultrasonic beam can be perpendicular to the arterial wall
even in the case of subjects with different distances from the skin surface to the artery.
2.4. Viscosity estimation by remote actuation
For assessment of viscosity, changes in blood pressure at multiple frequencies weregenerated by applying the external cyclic actuation from the skin surface on the brachialartery [22]. Resultant change in blood pressure propagates along the artery and causeschange in wall thickness at the carotid artery. In the case of sinusoidal actuation, Eq. (1)68

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can be rewritten as the complex elastic modulus by replacing Dp and Dqm,n,max by Dp 0e jxt
and De m,n,0e j(xt+u) as follows:
1 qm;n;0
Nm  n þ 1
Dp0 juþe ;ð5ÞE4h;m;n ðxÞ ¼
2
Nm
h0
Dqm;n;0where u is the phase difference between changes in pressure and wall thickness. Eq. (5)shows that the elastic modulus obtained from the amplitudes, Dp 0 and Dqm,n,0, of changes
in blood pressure and strain (change in thickness) corresponds to the absolute value,*|E h,m,n(x)|, of the complex elastic modulus.
When the Voigt model can be assumed as a viscoelastic model of the wall, the absolute* (x)|, of the complex elastic modulus is expressed by the static elastic
value, |E Voigt
modulus, E s, and viscosity constant, D, as follows:
qﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃ
4
ð
Þj
x
¼
Es2 þ ðxgÞ2 :jEVoigt
ð6Þ
From the in vitro experiments, using extracted human femoral arteries, it was shown that(xg)2/E 2s c10 [23]. Under such a condition, Eq. (6) can be approximated as follows:
E V4oigt ðxÞcxg:ð7Þ

From Eq. (7), the viscosity constant, D, can be obtained by estimating the slope of themeasured |E *h,m,n (x)|.

3. Experimental results
3.1. Measurement in the long-axis plane3.1.1. In vitro determination of statistical parameters
Immediately after nine iliac arteries (25–40 mm in length and 4–24 mm in outer diameter)with plaques extracted from patients with embolism, cross-sectional elasticity distribution,
E h,m,n , was measured using the above method (Fig. 2(a)) under the same artificial circulation
system to generate a change in pressure so that it ranged from the diastolic pressure to thesystolic pressure of the subjects. After each in vitro measurement, elastica–Masson stain wasapplied. (Two typical results are shown in Fig. 2(b).) From the stain images, 10 regions witheither lipid or a mixture of smooth muscle and collagen fiber were assigned in E u,m,n of thenine specimens. (Each histogram of E h,m,n in the respective regions is shown in Fig. 2(c).)
For the respective categories of the nine arteries, the average and the standard deviation inelasticity were determined to be 81F40 kPa and 1.0F0.63 MPa, which were registered as
the reference parameters (Fig. 2(d)).
Based on these reference parameters, each point in the cross-sectional elasticity
distribution, which had been noninvasively measured by the above method in separate invivo experiments, was statistically classified as one of three categories (lipid, mixture ofsmooth muscle and collagen fiber, or others). Thus, the arterial wall and the atherosclerotic
plaque were electronically stained [17].

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Fig. 2. By referring to the stained image (b), the regions of lipid and of the mixture of smooth muscle and collagenfiber were assigned in the spatial distribution of elasticity E h,m,n measured in in vitro experiments (a). From thehistogram of E h,m,n for each region (c), the elasticity for each category was registered as a reference (d), fromwhich each tissue was categorized as belonging to one of the three categories [17].

3.1.2. In vivo measurement at common carotid arteries
The proposed method was applied to in vivo measurements of the common carotid
arteries of a healthy subject (Fig. 3(a)) and a patient with hyperlipidemia having
atherosclerotic plaque (Fig. 3(b)). In Fig. 3, for each subject, a cross-sectional image
obtained by conventional ultrasound diagnostic equipment is shown at the top. The crosssectional elasticity distribution of E h,m,n was color-coded and superimposed on thereconstructed B-mode image as shown in the middle. Finally, at the bottom, thecategorized result is shown. The lipid and the mixture of smooth muscle and collagen fiber
are shown by yellow and cyan, respectively, and the category of botherQ is not colored.
3.2. In vivo measurement at common carotid artery in the short-axis plane
The change in the thickness of a common human carotid artery of a 29-year-old male
was measured in the short-axis plane. The B-mode image measured by the linear-typeprobe is shown in Fig. 4. The B-mode images were drawn at the time of the R-wave of theelectrocardiogram. Figs. 4(a) and (b) show B-mode images measured by the conventional
linear scanning method and the proposed method, respectively. In this experiment, L i =15.5mm and the largest beam angle, h max, was 21.28. In Fig. 4, beams at positions l 1 and l 2
measured the same regions of the posterior wall, as measured by beams at the positions, l 1Vand l 2V.

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Fig. 3. Common carotid arteries of (a) a healthy male subject (24-year-old) and (b) a male patient withhyperlipidemia (71-year-old). The area outlined by green in each uppermost figure was analyzed in the middle
and bottom of the figure [17].

Fig. 4. B-mode images measured by the conventional linear scanning method (a) and the proposed method (b).

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M-mode images, electrocardiograms, velocities at the intimal side and the adventitialside of the posterior wall, and changes in wall thickness measured at beam positions l 1, l 2,
l 1V and l 2V in Figs. 4(a) and (b) are shown in Figs. 5(a), (b), (c) and (d), respectively. Bylinear scanning, as shown in Fig. 5(a-5), the change in thickness was measured with highreproducibility at the beam position l 1, where the ultrasonic beam was perpendicular to thearterial wall. However, at the beam position l 2 (Fig. 5(b-5)), where the ultrasonic beam
was not perpendicular to the arterial wall, the amplitude of the change in thickness wasdifferent from that measured at l 1 because the direction of wall expansion due to theheartbeat slipped off the beam. Moreover, the reproducibility during heartbeats was madeworse by the decrease in the signal-to-noise ratio of the reflected ultrasonic waves. The
ratio, A(l 2)/A(l 1), of amplitudes A(l 1) and A(l 2) of echoes at the beam positions l 1 and l 2 inFig. 4(a) was 0.059.
As shown in Figs. 5(c-5) and (d-5), using the proposed method, the changes in the
thickness of the arterial wall caused by the heartbeat were measured with a similar
amplitude, waveform, and high reproducibility at both beam positions l 1V and l 2V. The ratio

Fig. 5. In vivo experimental results obtained by the linear scanning method at beam positions l 1 (a) and l 2 (b) andby the proposed method at beam positions l 1V (c) and l 2V (d). (1) M-mode image. (2) Electrocardiogram. (3)Velocity on intima. (4) Velocity on adventitia. (5) Change in thickness of the posterior wall.

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A(l 2V)/A(l 1V) was 0.92. At the beam position l 2V, a beam cannot be scanned to beperpendicular to the wall by the linear scanning method.
Figs. 6(a) and (b) shows the B-mode images of the carotid artery obtained byconventional linear scanning at two different angles. The regions shown by the yellow
dotted lines and red dotted lines in Fig. 6(a) correspond to those shown in Fig. 6(b). Figs.
6(c) and (d) shows the cross-sectional images of elastic modulus measured in two different
regions, shown by yellow dotted lines and red dotted lines in Figs. 6(a) and (b),respectively. Figs. 6(c) and (d) was measured at the same angles as Figs. 6(a) and (b),respectively. The elastic modulus is obtained by the maximum change in thickness and thepulse pressure measured at the upper arm. The resolution of the cross-sectional image ofelastic modulus was 80 Am in the depth direction and 1.438F0.58 in the circumferential
direction.
The region of the carotid artery between two red dotted lines in Fig. 6(b) was measured.
Then the probe was rotated about 208 along the neck, and the region between two yellow
dotted lines in Fig. 6(a) was measured. It is recognized that the elasticity images of theregions between two green dotted lines shown in Figs. 6(c) and (d), which are measured byboth measurements, correspond to each other. The mean value and the standard deviationof the measured elastic modulus shown in Figs. 6(c) and (d) were 0.18F0.14 and
0.2F0.15 MPa, respectively.
From these results, using the proposed method, the elasticity distribution of the arterial
wall can be measured transcutaneously within the wide region in the short-axis plane.
3.3. In vivo viscosity measurement at common carotid artery
Fig. 7(a-1) shows the M-mode image of a human carotid artery of a 30-year-old male.
Fig. 7(a-3) shows the blood pressure measured at the radial artery with an applanation
tonometer. By setting two points, A (intimal side) and B (adventitial side), along an
ultrasonic beam at the R-wave of electrocardiogram, the velocities, v A(t) and v B(t), of thesepoints were obtained by the phased tracking method as shown in Figs. 7(a-4) and (a-5).

Fig. 6. (a) and (b) are B-mode images of the carotid artery in the short-axis plane obtained by the conventional
linear scanning method. (c) Cross-sectional elasticity image measured from the direction shown with yellow lines
in (a). (d) Cross-sectional elasticity image measured from the direction shown with red lines in (b).

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Fig. 7. (a) (1) M-mode image. (2) Electrocardiogram. (3) Blood pressure. (4) Velocity at intimal side. (5) Velocity
at adventitial side. (b) Absolute value of complex elastic modulus plotted as a function of actuation frequency.

In Figs. 7(a-4) and (a-5), it is found that there are velocity components caused by remote
actuation at each actuation frequency.
In systole, there are velocity components caused by the heartbeat, and it is difficult todiscriminate components caused by remote actuation from those caused by the heartbeat.
Therefore, the amplitude of velocity components caused by remote actuation ismeasured in diastole by applying Fourier transform with a Hanning window at 0.5 s
after R-wave of the electrocardiogram. From the estimated amplitudes of the change ininternal pressure and the change in wall thickness, the absolute value of complex elastic
modulus at each actuation frequency was obtained based on Eq. (5) (in this experiment,
M=N m =1) and was plotted as a function of actuation frequency as shown in Fig. 7(b). In
Fig. 7(b), it is found that the absolute values of elastic modulus for three subjects increase
with actuation frequency. The viscosity constants were estimated to be 2.1 kPad s (30-yearold male), 6.5 kPad s (22-year-old male), and 1.6 kPad s (24-year-old male), respectively,
and the order of these values was the same as those reported in the literature (about 5
kPad s) [24].

4. Conclusion
In this study, the regional viscoelasticity of the arterial wall was measured withtranscutaneous ultrasound. This novel approach offers potential for the diagnosis of thevulnerability of plaque in a clinical setting.

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