Japanese Journal of Applied Physics
Vol. 46, No. 7B, 2007, pp. 4881–4888#2007 The Japan Society of Applied Physics

Ultrasonic Measurement of Change in Elasticity due to Endothelium Dependent
Relaxation Response by Accurate Detection of Artery-Wall Boundary
Takuya KANEKO, Hideyuki H ASEGAWA, and Hiroshi K ANAI
Graduate School of Engineering, Tohoku University, Sendai 980-8579, Japan(Received November 24, 2006; revised March 16, 2007; accepted April 25, 2007; published online July 26, 2007)

Ross hypothesized that an endothelial dysfunction is considered to be an initial step in atherosclerosis. Endothelial cells,
which release nitric oxide (NO) in response to shear stress from blood ﬂow, have a function of relaxing smooth muscle in the
media of the arterial wall. For the assessment of the endothelial function, there is a conventional method in which the change
in the diameter of the brachial artery caused by ﬂow-mediated dilation (FMD) is measured with ultrasound. However, despite
the fact that the collagen-rich hard adventitia does not respond to NO, the conventional method measures the change indiameter depending on the mechanical property of the entire wall including the adventitia. Therefore, we developed a methodof measuring the change in the thickness and the elasticity of the brachial artery during a cardiac cycle using the phased
tracking method for the evaluation of the mechanical property of only the intima-media region. In this study, the initial
positions of echoes from the lumen-intima and media-adventitia boundaries are determined using complex template matchingto accurately estimate the minute change in the thickness and the elasticity of the brachial and radial arteries. The ambiguity inthe determination of such boundaries was eliminated using complex template matching, and the change in elasticity measured
by the proposed method was larger than the change in inner diameter obtained by the conventional method.
[DOI: 10.1143/JJAP.46.4881]
KEYWORDS: complex template matching, elasticity, atherosclerosis, endothelial dysfunction, nitric oxide, flow-mediated dilation

1.

Introduction

The main cause of circulatory diseases is considered to beatherosclerosis. Ross hypothesized that endothelial dysfunction is considered to be the initial step in atherosclerosis.1)
Consequently, it is important to diagnose the endothelial
dysfunction for early preventive treatment.
Morphologically, the arterial wall consists of three
concentric layers called intima, media, and adventitia.2) In
normal muscular arteries, the tunica intima consists ofmono-layered endothelial cells. The media is composed
predominantly of smooth muscle cells. The adventitia,
which is composed of ﬁbrous connective tissue, covers theoutside of the intima-media complex. Endothelial cells
mainly adjust the distensibility of blood vessels and maintainvascular homeostasis.
Although there are techniques for evaluating the intimamedia thickness3,4) and elasticity of an arterial wall5–7) for
the diagnosis of atherosclerosis, ﬂow-mediated dilation(FMD) measurements in brachial and other conduit arteries
have become popular means of evaluating the endothelial
function for the early diagnosis of atherosclerosis.8,9)
Endothelial cells, which release nitric oxide (NO) in
response to shear stress from blood ﬂow,10) have a functionof relaxing smooth muscle in the media of the arterial wall.
In FMD measurement, the endothelium-dependent change in
inner diameter in response to reactive hyperemia ismeasured by comparing the luminal diameter of the brachialartery before and after the ischemia of the forearm induced
by pressurizing a cuﬀ. However, the collagen-rich hardadventitia does not respond to NO. Therefore, the conventional method underestimates the FMD caused by an
endothelial response because the change in inner diameter
depends on not only the mechanical properties of the intima
and media but also that of the adventitia.
Sugimoto et al. proposed a method of measuring theE-mail address: hkanai@ecei.tohoku.ac.jp

change in thickness and the elasticity of the brachial arterial
wall during a heart cycle using the phased tracking method
to evaluate the change in the mechanical property of only the
intima-media region.11–13) In this measurement, a method isnecessary to assign the boundaries of the intima-media layer
for objective assessment. If a conventional edge detector
such as the simple thresholding of magnitude is applied tothe ultrasonic RF signal, many fragmented edges in thearterial wall will be detected by the inﬂuences of scatterers
and interfaces, and the true media-adventitia boundary is noteasily identiﬁed because an echo from the media-adventitia
boundary includes redundant echoes from the intima-media
complex. Fun et al. proposed a technique for the boundary
detection of the brachial arterial wall using the templatematching between an envelope of the ultrasonic RF signal
and an adaptive approximate Gaussian model.14)
In this study, the lumen-intima boundary (LIB) andmedia-adventitia boundary (MAB) are detected using thecomplex template matching between the measured complex
demodulated signal and the adaptive complex model signal
to obtain transients in the change in thickness and theelasticity of the intima-media region. Additionally, theproposed method is applied to the measurement of conventional FMD, the change in intima-media thickness, and theelasticity of the radial artery using high-frequency ultrasound. FMD measurement in the radial artery, particularly
the measurements of the change in elasticity due to an
endothelium-dependent relaxation response, has not yet beenreported in the literature. There is an inversely proportional
relation between FMD value and the inner diameter of theartery at rest because ﬂow rate, which is related to shear
stress, is inversely proportional to the inner diameter.15)
Therefore, the radial artery would be a more suitable site forthe measurement of FMD.
2.

Principles

2.1 Elastic modulus estimation of arterial wall
To calculate a minute change in wall thickness caused by4881

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T. KANEKO et al.

c0 ^ðtÞv^ðtÞ ¼ ;
2!0 T(a)

2.5 mmdepth

5.0 mmlumen
LIB echolumen
MAB echoð2:1Þ
LIB echo

where !0 and c0 are the center angular frequency of theultrasound wave and the speed of sound, respectively. Thechange in thickness, hðtÞ, between two diﬀerent positionsin the arterial wall along an ultrasonic beam is obtained fromthe diﬀerence between displacements, xA ðtÞ and xB ðtÞ, atthese two positions ash^ðtÞ ¼ x^A ðtÞ  x^B ðtÞZt¼ fv^A ðtÞ  v^B ðtÞg dt:
ð2:2Þ

MAB echo

Fig. 1. Reconstructed B-mode images of (a) brachial and (b) radial
arteries of 33-year-old male in plane perpendicular to axis of artery.

0

where r0 and h0 are the internal diameter and wall thickness
at the end diastole, respectively.
Determination of optimum initial positions of arterywall boundaries
Figure 1 shows the cross-sectional ultrasonic images ofthe brachial and radial arteries. As shown in Fig. 1, along an
ultrasonic beam, the ultrasonic echo reﬂected from the LIBof the posterior wall can be found, the amplitude of the echoslightly decreases in the media region, and it increases again
around the MAB.16) In this study, echoes from LIB and
MAB of the posterior wall should be tracked to accurately
obtain the minute change in the intima-media thickness andelasticity. However, it is diﬃcult to detect the MAB of thearterial wall because weak echoes from the intima-media
region overlap with the echo from the MAB, and the leadingedge of the echo from the MAB is not clear. Therefore, theMAB cannot be determined directly from the B-mode image
shown in Fig. 1. Consequently, in this study, a method of
detecting the initial positions of these boundaries using thecomplex template matching between RF echoes and model
echoes is investigated, and is applied to the detection of the
LIB and MAB of the posterior wall.
Figure 2 shows a schematic view of ultrasonic echoesfrom the artery-wall boundaries. The ultrasonic echo signal
from the posterior wall includes mainly echoes from the LIB

US echo

US pulse

The change in thickness, hðtÞ, corresponds to the incremental strain in the radial direction at a time t due to apressure increment, pðtÞ, from the diastolic pressure.
Therefore, from the maximum change in wall thickness,hmax ¼ max jhðtÞj, and the pulse pressure pmax ¼
max pðtÞ, which is the diﬀerence between the systolicand diastolic pressures, the approximate circumferential
elastic modulus Eh is obtained as12)


1 r0
pmax
Eh 
þ1
;
ð2:3Þ
hmax
2 h0
h0

2.2

(b)

depth

a heartbeat, the velocity of the artery-wall boundary is
estimated.11) The velocity vðtÞ is estimated from the phaseshift ðtÞ of echoes in two consecutive frames. The phaseshift ðtÞ is obtained on the basis of the cross-correlation
applied to the demodulated signal zðt; d þ xðtÞÞ of RF echo
reﬂected at a depth d þ xðtÞ at a time t, where d and xðtÞ are
the initial depth set at t ¼ 0 and the displacement of thearterial wall in the direction of depth, respectively. From theestimated phase shift ^ðtÞ, the velocity of the arterial wall
during a pulse repetition interval T is obtained aslumen
intimamedia
adventitia

o

depth

x 1 (t)x 2 (t)Fig. 2. Schematic view of ultrasonic echo from vessel wall whenultrasonic beam passes through center of artery.

and MAB. Therefore, we modeled the echo from the arterial
wall by those from the LIB and MAB, x^1 ðtÞ and x^2 ðtÞ. The
model echoes x^1 ðtÞ and x^2 ðtÞ are simulated by referring to theultrasonic echo from a point scatterer (35-mm-thick ﬁne wirein water), as shown in Fig. 3. The model echoes x^1 ðtÞ andx^2 ðtÞ are formulated by multiplying the sinusoidal waves atthe center frequencies f1 and f2 by the Hanning window
wðtÞ, which shows the envelope of an echo, asx^1 ðtÞ ¼ a1  sin½2 f1 ðt  1 Þ  w1 ðt  1 Þ;
ð2:4Þ
ð2:5Þ
x^2 ðtÞ ¼ a2  sin½2 f2 ðt  2 Þ  w2 ðt  2 Þ;
2fi twi ðtÞ ¼ 0:5  0:5 cosði ¼ 1 or 2Þ; ð2:6Þ
nwhere f1 and f2 are the center frequencies of the LIB and
MAB echoes, a1 and a2 are amplitude coeﬃcients of echoesfrom the LIB and MAB, and 1 and 2 are time delays ofrespective echoes, respectively; n is the number of cycles atthe center frequency in an echo. As shown in Fig. 3, themodel echo is in very good agreement with the measured
one. The model echo x^ðtÞ is expressed by the sum of x^1 ðtÞand x^2 ðtÞ as4882

Jpn. J. Appl. Phys., Vol. 46, No. 7B (2007)(a) 1.2

T. KANEKO et al.

z^ðtÞ are the complex demodulated signals of xðtÞ and x^ðtÞ,respectively.
At the preassigned lags 1 and 2 , the normalized meansquare error (MSE)  between the measured RF and model
echoes is calculated. However, in the case that the MSE isnormalized throughout the period to calculate it, thenormalized MSE  becomes insensitive to the goodness ofﬁt with respect to x1 ðtÞ because the echo from the LIB ismuch smaller than that from the MAB. Therefore, thenormalized MSE  is separately calculated with respect tothe pulse durations of x^1 ðtÞ and x^2 ðtÞ. The region-to-region
normalized MSE  is formulated asXjzðmTs Þ  z^1 ðmTs Þj2

measurement signal
model signal

0.8

amplitude

0.4
0
-0.4
-0.8
-1.2

0

0.2

0.4

0.6

0.8

time [ µ s](b) 1.2

1

1.2

1.4

1.6

¼

measurement signal
model signal

m2K\K\LX

m2K\K\LX
jzðmTs Þ  z^2 ðmTs Þj2

0.8

amplitude

0.4

þm2L\L\KX

0jzðmTs Þj2

X m2L\L\KjzðmTs Þ  f^z1 ðmTs Þ þ z^2 ðmTs Þgj2
m2K\L
X
þ
;
jzðmTs Þj2

-0.4
-0.8
-1.2

jzðmTs Þj2

ð2:11Þ

m2K\L

0

0.1

0.2

0.3

0.4

0.5 0.6
time [ µ s]

0.7

0.8

0.9

1

Fig. 3. Measured RF signal [(a) 10 MHz, (b) 22 MHz] from a point
scatterer (dashed line) and model signal (solid line).

x^ðtÞ ¼ x^1 ðtÞ þ x^2 ðtÞ:

ð2:7ÞThe parameters f fi g, fai g, and fi g (i ¼ 1; 2) are determinedsuch that the diﬀerence between the model x^ðtÞ andmeasured echo xðtÞ is minimal. At ﬁrst, the instantaneous
center frequencies f fi g of the measured RF signal are
estimated using complex auto-correlation.17,18) An acquired
RF signal is denoted by xðmÞ, where m shows the time mTs(Ts : sampling period of echoes). The acquired RF signal isquadrature-demodulated by applyingxðmÞ ¼ LPF½xðmÞ  e j!dem mTs ;ð2:8Þ

where LPF½ represents low-pass ﬁltering, and fdem ¼ !dem =2 is the demodulation frequency. To estimate the diﬀerence
between the demodulation frequency fdem and the actual
center frequency f0 , the complex auto-correlation coeﬃcientRðm0 Þ at lag m ¼ 1 is calculated as
M
1
XR^ðmÞ ¼
ð2:9Þ
x ðmÞ  xðm þ mÞ;m¼0

where z^1 ðtÞ and z^2 ðtÞ, which correspond to x^1 ðtÞ and x^2 ðtÞ, arecomplex demodulated model signals of x^1 ðtÞ and x^2 ðtÞ withthe pulse lengths K and L, respectively. Figures 4(c) and4(d) show two cases in which x^1 ðtÞ and x^2 ðtÞ overlap and do
not overlap, respectively. The overlap of echoes from the
LIB and MAB is automatically taken into consideration
because the normalized MSE is evaluated between themeasured echo xðtÞ and model x^ðtÞ which is the sum of x1 ðtÞand x2 ðtÞ. It is shown by the third term on the right-hand sideof eq. (2.11) that the diﬀerence between the measured zðtÞand the sum of the models z^1 ðtÞ and z^2 ðtÞ is evaluated. The
phase information of the echo from the LIB and that from
MAB are considered because z^1 ðtÞ and z^2 ðtÞ are complex
signals. For the robust detection of the leading edge of anecho from the LIB, a constraint term, , is added by
calculating the variance in xðtÞ just before the LIB echo, asshown in Fig. 4. The variance  is normalized by the mean
power  2 in the region of interest before the the model LIB
echo x^1 ðtÞ. The magnitudes of echoes from the lumen are sosmall, and the normalized variance  is nearly zero. On theother hand,  increases when the model LIB echo x^1 ðtÞ islocated on the wall (false detection) because the region
shallower than x^1 ðtÞ is not the lumen in such a situation, andthere are echoes. Therefore, the number of cases of false
detection can be reduced by adding . The ﬁnal form of thenormalized MSE  is given by

where MTs is set to one wavelength of the ultrasonic signal.
= RðmÞ, the actual RF center frequencies,
From the phase f fi g (i ¼ 1; 2), can be estimated as18)f^i ¼ fdem = R^ ðmÞ

2Ts

ði ¼ 1; 2Þ: ¼  þ ;
¼

ð2:10Þ

The amplitude ratio between the measured RF echo xðtÞand the model echo x^ðtÞ, which is composed of x^1 ðtÞ and x^2 ðtÞ
obtained using the estimated center frequencies f fi g, iscalculated by the least-squares method. Note that zðtÞ and

X
1 =Ts

jxðmTs Þj2;
2
m¼1 =Ts N

ð2:12Þ
ð2:13Þ

where N is the number of samples necessary for calculating
the variance and is set to approximately three wavelengths.
According to the above process, the normalized MSE is
calculated with respect to each preassigned 1 and 2 . The
time delays 1 and 2 were changed at intervals of f=ð6!i Þg

4883

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T. KANEKO et al.

1.2

1.2

K

0.8

normalized amplitude

normalized amplitude

(a)

Ν

0.4
0
-0.4
-0.8

(c)

L

0.8
0.4
0
measured
RF signal
model signal of
the LIB echomodel signal of
the MAB echo

-0.4
-0.8

measured RF signal
-1.2

-1.2

depth

200 µ m

1.2

1.2
0.8
0.4
0
-0.4
model signal of
the LIB echomodel signal of
the MAB echo

-0.8
-1.2

200 µ m

depth
K

K K L
normalized amplitude

normalized amplitude(b)

200 µ m
Ν

L
L K L

0.4
0
measured
RF signal
model signal of
the LIB echomodel signal of
the MAB echo

-0.4
-0.8
-1.2

depth

(d)

0.8

200 µ mdepth

Fig. 4. (a) Measured RF signal. (b) Model signal. Periods for calculating MSE between measured RF signal and model signal in casesof (c) separated x^1 ðtÞ and x^2 ðtÞ and (d) overlapped x^1 ðtÞ and x^2 ðtÞ.

(i ¼ 1; 2). In this study, the diﬀerence 2  1 ranged from100 to 500 mm by referring to the typical intima-media
thickness.19,20) The time delays 1 and 2 that minimize thenormalized MSE between the measured RF signal andmodel signal are determined as the optimum boundary
positions k and l, respectively.
2.3 In vivo measurement
In this study, the right brachial and radial arteries of ahealthy subject (subject A: 33-year-old male) and the rightbrachial arteries of three healthy subjects (subjects B, C, D:
22-year-old males) are measured. In the measurement of thebrachial artery, the RF echoes were acquired using a
10 MHz linear-type ultrasonic probe of ultrasonic diagnostic
equipment (SSD-6500, Aloka). The received echoes are
acquired at a sampling frequency of 40 MHz. At ﬁrst, wemeasured the electrocardiogram and RF echoes during two
consecutive heartbeats for 2 min at intervals of about 20 s at
rest. Then, blood ﬂow is stanched by surrounding theforearm with a cuﬀ at a pressure of 250 mmHg for 5 min.
After the release of the cuﬀ, NO, which is generated by theendothelium in response to the shear stress due to increased
blood ﬂow, makes the smooth muscle in the media to relax.
RF echoes are acquired for 3 min at an interval of about 12 s
after the release of the cuﬀ. For the measurement of thesame position where the intima-media region of the brachialposterior wall is clearly imaged, the ultrasonic probe washeld by the pointsetter. Blood pressure was continuously
measured at the wrist with a spygnomanometer, and thesystolic and diastolic blood pressure in each measurement
were recorded.
For the measurement of the radial artery, sets of RF
echoes were acquired using a 22-MHz linear-type ultrasonicprobe of another modiﬁed ultrasonic diagnostic equipment(Diasus, Dynamic Imaging). The received echoes are
acquired at a sampling frequency of 66.5 MHz using acustom-made system. The measurement procedure is similar
to that for the brachial artery described above.
3.

In vivo Experimental Results

3.1 Example of complex template matching
Figure 5 shows the results of applying the complex template matching method to a measured RF echo. Figures 5(a)and 5(b) show the normalized MSE distributions 20 log10 
and 20 log10  without and with a constraint term, respectively. As shown in Fig. 5(a), the MSE without  converges
at several combinations of 1 and 2 . In contrast, MSE
converges at one point in accordance with the constraint, asshown in Fig. 5(b).
The RF signals and envelopes of the measured signal andthe echo model at point A in Fig. 5(a) are shown in Fig. 6.
The estimated frequencies of the LIB and MAB echo models
are respectively 17.57 and 17.17 MHz, and the echo models
are in good agreement with the measured signal. The model
can represent the measured signal very well because theMSE  at point A in Fig. 5(a) is 12%.
Figure 7 shows the result of the boundary detection for theseveral ultrasonic beams in a B-mode image along the4884

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T. KANEKO et al.

(a)power [dB]B

200 µ m0 µm∼

A100 µ m ∼200 µ mτ2

power [dB]
0 µm ∼

200 µ m

100 µ m ∼200 µ m

τ2

(b)

lumen

k
l

Fig. 7. (a) Reconstructed B-mode image in long-axis view of radial arteryof 33-year-old male. (b) Result of boundary detection using complex
template matching.

40
30
20
10
0
-10
-20

τ1

l

2.5 mm

0
-5
-10
-15

(b)

k

(a)

30
20
10
0
-10
-20

τ1

depth
(a)

2.5 mm

k
l
(b)

ECG
1 mm/s

(c)

0

0
-5
-10
-15

- 1 mm/s
1 mm/s(d)0- 1 mm/s
10 µ mFig. 5. Results of applying complex template matching to RF data. (a)20 log10  vs time delays 1 and 2 . (b) 20 log10  vs time delays 1 and 2 .

(e)

0
- 10 µ m

amplitude

0
1.2
0.8
0.4
0
-0.4
-0.8
-1.2

0.2

0.4

0.6

2

0.8

1

1.2

measurement signal
LIB echo model
MAB echo model

1.6
1.2
0.8
0.4
0

0

0.2

0.4

0.6

time [ µ s]

0.8

1

1

1.5

2

2.5

time [s]Fig. 8. (a) M-mode image of brachial artery (healthy 33-year-old male).
(b) Electrocardiogram. (c) Velocity at point k assigned as LIB. (d)Velocity at point l assigned as MAB. (e) Change in intima-media
thickness.

measurement signal
LIB echo model
MAB echo model

0

0.5

1.2

Fig. 6. RF signal and envelope of measured signal and echo models.

longitudinal direction of the radial artery. As shown in
Fig. 7, the boundary positions, which are shown as k and l on
the B-mode image, are detected appropriately and objectively by complex template matching.
3.2 Measurement in brachial artery
Figure 8 shows an example of measurement of the change
in intima-media thickness in the healthy 33-year-old maleduring two consecutive heartbeats for 3 s at rest. As shown in
Fig. 8(a), the initial positions k and l of the lumen-intima
and media-adventitia boundaries of the brachial artery on the
M-mode image were determined by the method described in§2.2, and then the instantaneous positions of these points
were automatically tracked as shown by black lines using thephased tracking method. Figures 8(c) and 8(d) show theestimated velocities at k and l, respectively. The change inthe thickness of the intima-media complex was calculated by
the temporal integration of the diﬀerence between thesevelocities, as shown in Fig. 8(e). The elastic modulus of theintima-media region of the brachial artery for each measurement was calculated using the estimated maximum change
in thickness and the pulse pressure, as described in §2.1.
Figure 9 shows the transients in the maximum change inthe thickness of the intima-media region of the brachialarterial wall during one cardiac cycle and the pulse pressurein each measurement. The change in thickness and the pulsepressure at each RF acquisition during 10 min are plotted bycrosses and diamonds, respectively. About 20 to 40 s after
the release of the cuﬀ, the change in thickness increased
compared with that at rest, as shown by the dashed line in
Fig. 9. Then, the change in thickness gradually returned tothat at the rest.
Figure 10 shows the transients in the change in innerdiameter manually obtained from the B-mode image at eachR-wave of the electrocardiogram and the change inelasticity, which is calculated by the maximum change in4885

Jpn. J. Appl. Phys., Vol. 46, No. 7B (2007)

T. KANEKO et al.

70

17

60

11

50

8

40

5

30

change in thickness [ µm]

14

2

0
rest

0
80 120avascularization
300 s

40

20

40

60

80

pulse pressure [mmHg]

change in thickness
pulse pressure

20
100 120 140 160 180 200

release of the cuff

time [s]Fig. 9. Transient changes in maximum change in thickness of intima-media region of brachial artery and pulse pressure (healthy 33year-old male).

2600

4.7

2200

4.6

1800

4.5

1400

4.4

1000

4.3

600

0
rest

40

80

120

avascularization
300 s

0

20

40

60

inner diameter [mm]elasticity [kPa]elasticity
inner diameter

4.2
100 120 140 160 180 200

80

release of the cuff

time [s]Fig. 10. Transient changes in elasticity of intima-media region of brachial artery and inner diameter (healthy 33-year-old male).

the thickness of the intima-media region in each measurement. The inner diameter and change in elasticity at each RF
acquisition over 10 min are plotted by diamonds and crosses,
respectively. From about 20 to 40 s, the maximum decreasein elasticity shown by the dashed line compared with theaveraged elasticity at rest was 29.3%. The inner diametershown by the dotted line increased by 3.0% compared withthe averaged inner diameter at rest about 70 s after the
release of the cuﬀ. These results are comparable with thosepreviously measured for the brachial artery.8,13,15)
3.3 Measurement of radial artery
Figure 11(a) shows an M-mode image of the radial arteryof the same subject shown in Fig. 8. Figures 11(c) and 11(d)show the estimated velocities at points k and l, respectively.
The change in the thickness of the intima-media complex
was calculated by temporal integration of the diﬀerence
between these velocities, as shown in Fig. 11(e). The
subcutaneous tissue on the radial artery is less than that on
the brachial artery as shown in Fig. 1.
Figure 12 shows the maximum change in thickness of theintima-media region of the radial artery estimated at ﬁveneighboring ultrasonic beams during one cardiac cycle andthe pulse pressure in each measurement. The averages of the

depth
(a)

2.5 mm

k
l

(b)
ECG
2 mm/s(c)0- 2 mm/s
2 mm/s(d)0- 2 mm/s
20 µ m(e)

0- 20 µ m
0

1

2

3

time [s]Fig. 11. (a) M-mode image of radial artery (healthy 33-year-old male). (b)Electrocardiogram. (c) Velocity at point k assigned as LIB. (d) Velocity atpoint l assigned as MAB. (e) Change in intima-media thickness.

maximum changes in the intima-media thickness in ﬁve
contiguous beams are plotted by crosses, and the verticalbars show the standard deviations. About 40 to 60 s after the
release of the cuﬀ, the maximum change in the thickness4886

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T. KANEKO et al.

70

30

24

60

18

50

12

40

6

30

0

0
rest

40

80 1200
avascularization
300 s

pulse pressure [mmHg]

change in thickness [ µ m]change in thickness
pulse pressure

20
100 120 140 160 180 200
time [s]20 40 60 80
release of the cuff

Fig. 12. Transient changes in maximum change in thickness of intima-media region of radial artery and pulse pressure (healthy 33year-old male).

1600

3.0

1200

2.5

1000

2.0

800

1.5

600

1.0

400

0
rest

40

80

120

avascularization
300 s

0

20

40

60

inner diameter [mm]1400
elasticity [kPa]3.5

elasticity
inner diameter

0.5
100 120 140 160 180 200

80

release of the cuff

time [s]Fig. 13. Transient changes in elasticity of intima-media region of radial artery and inner diameter (healthy 33-year-old male).

becomes maximum as well the measurement for the brachialartery.
Figure 13 shows the transients in inner diameter andelasticity, which is calculated from the average of the
maximum changes in thickness at ﬁve beams in eachmeasurement. About 40 to 70 s after the release of the cuﬀ,
the maximum decrease in elasticity compared with theaveraged elasticity at rest was 43.0%. The maximum
increase in inner diameter was 10.1% compared with theaveraged inner diameter at rest about 90 s after the release of
the cuﬀ. These results are comparable with those measured
for the brachial artery shown in Figs. 9 and 10.

maximum decrease in elasticity [%]0subject A (brachial)
subject A (radial)
subject B
subject C
subject D

–10
–20
–30
–40
–50
–60
–70
–80

0

2

4

6

8

10

12

14

maximum increase in diameter [%]3.4

Relationship between maximum increase in innerdiameter and maximum decrease in elasticity
The brachial arteries of subjects B, C, and D as well asthat of subject A are measured. The relationship between
the maximum increase in inner diameter and the maximum
decrease in elasticity for each subject is shown in Fig. 14. As
shown in Fig. 14, a negative correlation between the
maximum increase in inner diameter and the maximum
decrease in elasticity is conﬁrmed; it was found that the rateof change obtained by the proposed method is much larger
than that obtained by the conventional method. Moreover,Fig. 14. Relationship between maximum increase in inner diameter andmaximum decrease in elasticity for each subject.

the rate of change in the measurement of the radial artery ismuch larger than that of the brachial artery as shown by the
results of subject A.
4.

Discussion

In §3.4, the results of the measurement of the endothelium-dependent relaxation response of the brachial and4887

Jpn. J. Appl. Phys., Vol. 46, No. 7B (2007)

T. KANEKO et al.

maximum increase in diameter [%]12

and radial arteries due to endothelium-dependent ﬂowmediated dilation was measured. The normalized change inelasticity measured by the proposed method was much larger
than the normalized change in diameter measured byconventional methods, which evaluate the mechanicalproperty of the entire wall. Moreover, it is supposed thatthe change in the inner diameter and the elasticity of theradial artery are larger than those of the brachial artery
because the radial artery is rich in smooth muscle cells
which respond to NO, and the magnitude of the shear stressto the arterial wall is also larger. From these results, wedemonstrated the possibility of the sensitive assessment ofendothelial dysfunction at the radial artery using highaccuracy boundary detection.

subject A (brachial)
subject A (radial)
subject B
subject C
subject D

Pyke [15]
10
8
6
4
2
0
2.5

3

3.5

4

4.5

5

baseline diameter [mm]Fig. 15. Relationship between inner diameter at rest and maximum
increase in inner diameter.

radial arteries are shown. As shown in Fig. 14, comparedwith that for the brachial artery, the % maximum increase in
inner diameter and % decrease in the elasticity of the radial
artery were large. On the other hand, the rates of change inthe measurement of the identical subject had a characteristic
variation. Thus, we tried to explain the cause of the above
result.
The arterial wall consists of elastic ﬁbers, collagenous
ﬁbers and smooth muscle. The contents, sequences, anddistribution of these components diﬀer depending on thetype of artery and age.21) In particular, elastic arteries such as
the aorta are rich in elastic ﬁbers, and muscular arteries suchas the brachial and radial arteries, are rich in smooth muscle,
which actively extends the artery to adjust blood ﬂow. In
systemic circulation, the amount of elastic ﬁbers and smoothmuscle increases from the aorta to the peripheral vessel.22)
Therefore, it is considered that the rate of FMD of the radial
artery is much larger than that of the brachial artery because
the radial artery, which is more peripheral, is rich in smoothmuscle cells responding to NO.
Another reason is the diﬀerence in the magnitude of theshear stress exerted on the arterial wall due to blood ﬂow.
Shear stress exerted on the radial arterial wall is larger due tothe smaller diameter of the artery because a smaller vascular
diameter signiﬁcantly increases ﬂow rate.15) The relationshipbetween the inner diameter at rest and the maximum change
in the inner diameter of each subject is shown in Fig. 15. As
shown in Fig. 15, the characteristics of the rates of change inthe measurement of the brachial artery agree with those
provided by Pyke et al.15) Therefore, it is important tomeasure the FMDs in the arteries with similar diameters tocompare diﬀerent subjects.
5.

Conclusions

In this study, we signiﬁcantly reduced the ambiguity inassigning the lumen-intima and media-adventitia boundaries
using complex template matching. The transient change inelasticity only for the intima-media region of the brachial

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