Reprinted from

REGULAR PAPER

Noninvasive Measurement of Transient Change in Viscoelasticity Due to Flow-MediatedDilation Using Automated Detection of Arterial Wall Boundaries
Kazuki Ikeshita, Hideyuki Hasegawa, and Hiroshi Kanai

Jpn. J. Appl. Phys. 50 (2011) 07HF08# 2011 The Japan Society of Applied Physics

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REGULAR PAPER

Japanese Journal of Applied Physics 50 (2011) 07HF08
DOI: 10.1143/JJAP.50.07HF08

Noninvasive Measurement of Transient Change in Viscoelasticity Due to Flow-MediatedDilation Using Automated Detection of Arterial Wall Boundaries
Kazuki Ikeshita1 , Hideyuki Hasegawa1;2 , and Hiroshi Kanai2;1
12Graduate School of Biomedical Engineering, Tohoku University, Sendai 980-8579, Japan
Graduate School of Engineering, Tohoku University, Sendai 980-8579, Japan

Received December 13, 2010; accepted March 7, 2011; published online July 20, 2011
We measured the stress–strain relationship of the radial arterial wall during a heartbeat noninvasively. In our previous study, the viscoelasticity of
the intima–media region was estimated from the stress–strain relationship, and the transient change in viscoelasticity due to flow-mediateddilation (FMD) was estimated. In this estimation, it is necessary to detect the lumen–intima boundary (LIB) and the media–adventitia boundary(MAB). To decrease the operator dependence, in the present study, a method is proposed for automatic and objective boundary detection basedon template matching between the measured and adaptive model ultrasonic signals. Using this method, arterial wall boundaries wereappropriately detected in in vivo experiments. Furthermore, the transient change in viscoelasticity estimated from the stress–strain relationshipwas similar to that obtained manually. These results show the feasibility of the proposed method for automatic boundary detection enabling an
objective and appropriate analysis of the transient change in viscoelasticity due to FMD. # 2011 The Japan Society of Applied Physics

1. Introduction

The quantitative assessment of atherosclerosis, which is themain cause of circulatory diseases, is essential for making an
early diagnosis of such diseases. In early-stage atherosclerosis, endothelial dysfunction occurs.1,2) Additionally, it
has been reported that the characteristics of the smoothmuscle that composes the media of the artery change owing
to atherosclerosis.3) Consequently, it is important for an
early preventive treatment to noninvasively assess theendothelial function and mechanical properties of the media
mainly composed of the smooth muscle.
The intima, the innermost layer, is composed of endothelial cells and an internal elastic membrane. The endothelial
cells produce the endocrine substances for maintaining thehomeostasis of the vascular system.4) The smooth muscle
that composes the media is the main source of theviscoelasticity of the vessel wall. Therefore, the dilationand contraction of the artery depend on the characteristics ofthe media. Endothelial cells react to the shear stress causedby blood ﬂow and produce nitric oxide (NO), which is known
as a vasodepressor. The smooth muscle in the media isrelaxed as a result of the response to the produced NO.
For the evaluation of the endothelial function, a conventional technique is used to measure the transient change in
the inner diameter of the brachial artery caused by ﬂowmediated dilation (FMD) after the release of avascularization.5–9) However, this method cannot directly evaluate themechanical properties of the arterial wall. Additionally, theincrease in the inner diameter of the brachial artery is verysmall (about 0.4 mm) compared with the spatial resolution of
conventional ultrasonic diagnostic equipment of 0.1 mm.10)
Therefore, we have developed a method of directly measuring the changes in the elasticities of the intima–media
regions of the brachial and radial arteries during FMD.11,12)
Furthermore, we have simultaneously measured the stressand strain waveforms noninvasively, and the stress–strain
relationship during a heartbeat was identiﬁed.13) In ourprevious study, from the stress–strain relationship during aheartbeat, the viscoelasticity of the intima–media region wasResearch Fellow of the Japan Society for the Promotion of Science.
E-mail address: ikeshita@us.ecei.tohoku.ac.jp

estimated using the least-squares method, and the transientchange in viscoelasticity due to FMD was estimated.14)
In our method, it is necessary to measure the strain waveform using the phased tracking method.15) In this estimationof strain, the lumen–intima boundary (LIB) and media–
adventitia boundary (MAB) in an ultrasonic data must be
detected. To reduce the operator dependence, it is necessary
to detect the LIB and MAB positions objectively on themeasured radio frequency (RF) signals. In our previous
report, a complex method of template matching between themeasured ultrasonic RF signal and the adaptive model signal
is proposed for the automatic and objective boundary
detection of the radial artery.11) In this study, we improve
the adaptive model and conﬁrm the eﬃciency of this methodby a basic experiment using a rubber sheet. Additionally,
the transient changes in the stress–strain relationship andviscoelasticity of the arterial wall should be measured stably
for ten minutes. Although the method was previously
applied to the longitudinal scan of the arterial wall, it wasdiﬃcult to maintain an ultrasonic probe in place for ten
minutes; thus the scan plane passed through the central axisof the artery. In contrast, there should be a few ultrasonicbeams that pass through the central axis of the artery when
the artery is measured by a cross-sectional scan. Therefore,
for a stable measurement, in this study, the viscoelasticity
was measured by a cross-sectional scan of the artery.
2. Principles and Experimental Methods
2.1 Determination of optimum initial positions of artery–wall boundaries

Along an ultrasonic beam around the posterior wall, twodominant ultrasonic echoes are reﬂected from the LIB and
MAB. Ideally, the LIB and MAB are determined by referringto these echoes. However, the MAB is not easily identiﬁed
because an echo from the MAB includes redundant echoesfrom the intima–media complex. Therefore, if a simple edge
detector, such as a thresholding device, is applied to theultrasonic RF signals, misdetection frequently occurs owing
to the eﬀects of such redundant echoes.
To overcome the disadvantages of conventional gradientmethods, Fan et al. proposed a technique for the boundary
detection of the brachial arterial wall by template matching
between an envelope of an ultrasonic RF signal and an

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(a)

0

(b)

0.3

0.6

0.9

[µs]

anterior
wall

lumen

posteriorwall
at restafter avascularization

Fig. 1. (Color online) (a) RF echoes from posterior wall of radial artery and (b) the transient change in the RF echoes due to ﬂow-mediated dilation.

z^i ðnTs Þ ¼ Ci  exp½ j2f0 ðnTs   i Þ  !ðnTs  i Þ;82f0 nTs<
0:5  0:5 cos;!ðnTs Þ ¼
ði ¼ 1; 2; 3Þ
N
:
0;

ð1Þ
ð2Þ

1

measurement signal
model signal

0.8
Normarized amplitudeadaptive approximate Gaussian model.16) However, thismethod detected only the MAB because echoes from the
LIB are not always obtained by ultrasonic measurements. In
the present study, echoes from the LIB and MAB of theposterior wall should be tracked to accurately obtain theminute change in intima–media thickness. Therefore, it isnecessary to detect these boundaries appropriately andobjectively. Consequently, in this study, the initial positionsof these two boundaries were detected by template matching
between the measured RF echoes and the adaptive echomodel (not envelope).
Unlike a deterministic signal embedded in a noisy
background, the ultrasonic signal from the vessel wall
diﬀers from subject to subject, and the waveform of an echodepends on the center frequency and band width of an
ultrasonic system. Therefore, a ﬁxed template is not suitable.
Figure 1(a) shows RF echoes from the posterior wall of theradial artery, and Fig. 1(b) shows the transient change in
the RF echoes due to ﬂow-mediated dilation (FMD). The
boundary positions were detected manually on each RF
signal, and the RF signals were shown to align the detected
LIBs. We refer to these RF echoes to obtain an adaptive
template. Echoes from the posterior wall, including
redundant echoes from the intima–media complex, aremodeled by the sum of three models of a single RF echo.
A single RF echo is modeled by multiplying a sinusoidal
wave at the center frequency f0 by a Hanning window wðtÞ,which shows the envelope of an echo, as follows:

0.6
0.4
0.2
0
−0.2
−0.4
−0.6
−0.8
−1

0

0.1

0.2

0.3

0.4

0.5
[µs]

Fig. 2. (Color online) Measured RF echo signal from a point scatterer
and model signal.

where fCi g is the amplitude coeﬃcient of the echo, f i g is thetime delay of the echo, and n is the number of cycles at thecenter frequency in an ultrasonic pulse. To obtain complex
signals, quadrature demodulation was applied to themeasured signal. As shown in Fig. 2, the model echo is invery good agreement with the measured one. The proposed
adaptive template echo model z^ðtÞ is expressed by the linear
sum combination of f^zi ðtÞg (i ¼ 1; 2; 3), which respectively
model echoes reﬂected from the LIB, the intima–media
complex, and the MAB, as follows:z^ðtÞ ¼ z^1 ðtÞ þ z^2 ðtÞ þ z^3 ðtÞ:ð3ÞFirst, the center frequency f0 of each echo is roughly
estimated using the discrete Fourier transform, and then it isdetermined using complex autocorrelation.17) The parameters fCi g and f i g (i ¼ 1; 2; 3) are determined to obtain a

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(a)

(b)

2.5 mmgelatin gel(containing graphite)

2.5 mmultrasonic probe

0.5 mmrubbermodel signal

Fig. 3. (Color online) (a) Diagram and (b) B-mode image of the basic
experiment.

(b)0minimal squared diﬀerence between the model z^ðtÞ andmeasured echo zðtÞ. The normalized mean squared error(MSE)  is given by1
X
1NjzðnTs Þ  z^ðnTs Þj2
N n¼0
¼
;
1
X
1N
2
jzðnTs Þj
N n¼0

0.3

0.6

0

0.3

0.6

(c)

[µs]
0.9
1.2
redundant echo model signal
anterior boundary echo model signal
posterior boundary echo model signal

0ð4Þ

where N is the number of samples used for calculating ,
and Ts is the sampling interval. According to the above
process, the normalized MSE  was calculated with respect
to each preassigned f i g. The time delay f i g was changed atintervals of Ts =4.
Let us describe how the proposed method works.
Figures 3(a) and 3(b) show a diagram and a B-mode imageof the basic experiment using a rubber sheet. The thicknessof the rubber sheet is 0.5 mm, and it includes no scatterers.
On the other hand, the gelatin gel containing graphite hassome reﬂection from inside. In this experiment, the reﬂectedcomplex echo from the anterior boundary, from the gelatin
gel to the rubber sheet, imitated the echo from the MAB,
including a redundant echo from the intima–media complex,
and the single echo from the posterior boundary, from rubber
to air, imitated the echo from the LIB. Although the order ofthese three echo signals is diﬀerent from that in in vivo
measurements, the reﬂected echoes from these two boundaries were estimated to conﬁrm the eﬃciency of the templatematching method.
In this estimation, the echo from the anterior boundary
includes the redundant echoes from the gelatin gel. Therefore, in this basic experiment, the model signals arenumbered as the redundant echo (z1 ), anterior boundary(z2 ), and posterior boundary (z3 ). The time delay fi g(i ¼ 1; 2; 3) that minimizes the normalized MSE between themeasured RF signal and the model signal is determined. In
this estimation, anterior and posterior boundaries correspondto the MAB and LIB, respectively. Additionally, theamplitude of the MAB echo is larger than that of the LIB
echo in the in vivo measurement [for example, Fig. 1(a)].
Therefore, to avoid false boundary detection, the amplitude
coeﬃcient C2 is limited so that C1 < C2 and C3 < C2 in theminimization of the MSE. Additionally, the thickness
between the anterior and posterior boundaries ranged from
400 to 650 m. Figure 4(a) shows examples of the detectedboundaries on the B-mode image. Figure 4(b) shows themeasured echo and a model with a minimal . In Fig. 4(c),a model of each single echo is shown. The mean of the

0

0.9

1.2

[µs]

Fig. 4. (Color online) (a) Estimated boundaries on B-mode image.
(b) Measured and model signal z^. (c) Model of each single echo z^i .

estimated thicknesses (corresponding to  3  2 ) for ultrasonic beam positions is 453 m. Additionally, the model
echo is in very good agreement with the measured one[normalized MSE is 15.7% in Fig. 4(b)].
In this in vivo study, model signals are denoted as the LIB(z1 ), redundant echo (z2 ), and MAB (z3 ), and the distance
between the LIB and the MAB (corresponding to 3   1 )ranged from 100 to 400 m, as determined by referring to
the typical intima–media thickness of the arterial wall.18,19)
Additionally, to avoid false boundary detection, theamplitude coeﬃcient C2 is limited to C1 < C3 and C2 <C3 in the in vivo measurement, because the amplitude of the
MAB echo is much larger than those of the LIB andredundant echoes. The time delay fi g (i ¼ 1; 2; 3) thatminimizes the normalized MSE between the measured RF
signal and the model signal indicates the optimum positionsof the LIB, the redundant echoes from the intima–media
complex, and the MAB. The redundant echo was included in
the model to avoid the position of the redundant echo frombeing determined as the position of the MAB.
2.2 Estimation of artery–wall viscoelasticity using theleast-squares method

The detailed analysis of the change in the viscoelasticity of
the arterial wall due to FMD requires the in vivo measurement of the stress–strain relationship, which has not beenmeasured noninvasively thus far.
To determine the stress–strain relationship, the minutechange in the thickness (radial strain) hðtÞ of the rightradial arterial wall during a cardiac cycle was measured
using the phased tracking method.15) Together with themeasurement of ultrasonic RF signals for the estimationof hðtÞ, the waveform of blood pressure (stress) pðtÞ inthe left radial artery was continuously measured with asphygmometer.

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4

8

[mm]
12

(b)
0
depth

4

8

12

16

20

[mm]
24

2.5 mm(a) 0

2.5 mmdepth

K. Ikeshita et al.

Fig. 5. (Color online) B-mode images of (a) cross-sectional view, and (b) longitudinal view.

To obtain the change in thickness, the velocities of artery–wall boundaries (namely, the LIB and MAB) were estimated.
c of
The velocity vðtÞ was estimated from the phase shift ðtÞcechoes in two consecutive frames. The phase shift ðtÞ was
obtained using the complex cross-correlation applied to thequadrature demodulated signals of the measured RF echoes.
The change in thickness, hðtÞ, between two diﬀerent
depths, A and B (corresponding to the LIB and MAB,
respectively), in the arterial wall along an ultrasonic beam
was obtained from the diﬀerence between the displacements
xLIB ðtÞ and xMAB ðtÞ at these two positions as follows:^ ¼ x^LIB ðtÞ  x^MAB ðtÞhðtÞZt¼ ½v^LIB ðtÞ  v^MAB ðtÞ dt:
ð5Þ
0

The strain ðtÞ is obtained from the change in thickness,
which is divided by the intima–media thickness at the time of
the R-wave of the electrocardiogram.
By assuming the Voigt model as a viscoelastic model ofthe intima–media region, the stress–strain relationship isgiven by_ þ 0 ;^ ðtÞ ¼ Es ðtÞ þ ðtÞ

ð6Þ

where ^ ðtÞ is the stress modeled by the Voigt model, and_ðtÞ, ðtÞ,
Es , and  are the strain, strain rate, static elasticity,and viscosity, respectively. In the in vivo measurement, themeasured stress ðtÞ is the incremental strain due to the pulsepressure, whereas the measured stress includes the bias
stress (diastolic blood pressure). Therefore,  0 is added to theright-hand side of eq. (6) as the bias stress corresponding to
the diastolic pressure.
The parameters in eq. (6), namely, Es , , and  0 , areestimated using the least-squares method by minimizing themean squared error  between the measured ðtÞ and model^ ðtÞ stresses, which is deﬁned by
2

 ¼ Et f½ðtÞ  ^ ðtÞ g;

ð7Þ

where Et ½ indicates the averaging operation during a
cardiac cycle. The parameters E^ s , ^ , and ^ 0 that minimize 
are determined by setting the partial derivatives of  withrespect to Es , , and  0 to zero. To solve the resultant
simultaneous equations, the optimum parameters thatminimize  are determined.14)
2.3 Procedure for in vivo measurements

First, the proposed method was applied to the radial,brachial, and carotid arteries to show its feasibility inin vivo measurements. In these measurements, ultrasonic RF

echoes (transmit center frequency: 22 MHz) were acquired ata sampling frequency of 66.5 MHz for 2 s at a frame rate ofabout 160 Hz.
Moreover, the right radial artery of a healthy male subject(35 years old) was measured. Together with the measurement of RF signals, the waveform of blood pressure pðtÞin the left radial artery was continuously measured with asphygmometer (Colin Jentow-7700) for the estimation ofviscoelasticity.13,14) Before the measurement of the transientchange in viscoelasticity, the strain and blood pressure
waveforms were measured every 1 min for 10 min at rest toevaluate reproducibility. Subsequently, for the measurementof FMD, the measurements of strain and blood pressure wererepeated every 20 s for 2 min at rest before avascularizationand every 12 s for 3 min after recirculation.
For stable in vivo measurements, the radial artery wasmeasured in the cross-sectional view. Figures 5(a) and 5(b)show the B-mode images in the cross-sectional andlongitudinal views, respectively. In the cross-sectional view,the LIB can be recognized in ﬁve ultrasonic beams(corresponding to a lateral width of about 1 mm). This
indicates that the acceptable range of the movement in the
direction perpendicular to the arterial longitudinal directionis 1 mm. In the FMD measurement, it is diﬃcult to reducethe movement of an ultrasonic probe or arm to less than
1 mm for 10 min. In the longitudinal scan, a movement in theelevational direction larger than 1 mm makes it impossible toobtain echoes from the LIB at all beam positions. On theother hand, as described above, echoes from the LIB can beobtained at ﬁve beams even when the probe or arm moves
by more than 1 mm. Therefore, the cross-sectional view issuitable for the measurement of FMD.
3. Results
3.1 Example of template matching for boundary detectionof arterial wall

For the validation of the method, we applied templatematching to the radial, brachial, and carotid arteries.
Figures 6(a), 6(c), and 6(e) show the results of the boundary
detection for the radial, brachial, and carotid arteries,
respectively. As shown in Fig. 6, the LIB and MAB, which
are shown on the B-mode image, are detected appropriately
and objectively by the proposed template matching method.
Examples of RF signals and the determined echo models areshown in Figs. 6(b), 6(d), and 6(f). The mean squared errorsof these results were 13.8, 24.2, and 21.0%, respectively.
The echo models are in good agreement with the measured
signals.

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K. Ikeshita et al.

4

8

(b)

[mm]

2.5 mm

0

0

(c)

0

4

8

0.3

0.6

0.9 [µs]

0.3

0.6

0.9 [µs]

0.3

0.6

0.9

(d)

[mm]

2.5 mm

0

0
(e) 0

4

8

12 [mm](f)

2.5 mm

0

0

1.2
[µs]

Fig. 6. (Color online) Results of the template matching. (a, c, e) Detected boundaries shown on the B-mode images of the radial, brachial, and carotid
arteries, respectively. (b, d, f) Measured RF and model echoes.

3.2 In vivo experimental results for healthy subjects

evaluation of reproducibility at rest
0

2

4

4

8

[mm]
(b)

0

4

8

[mm]
(d)

6

(b)

(c)

(d)

8

[min]

10

0

4

8

[mm]

0

4

8

[mm]

2.5 mm

2.5 mm

2.5 mm
(c)

(a)

0
(a)

2.5 mmFor the evaluation of reproducibility, RF data and the bloodpressure waveform at rest were measured every 1 min for
10 min before the measurement of FMD.
Figure 7 shows examples of boundary detection of aradial artery. The LIB and MAB are detected appropriately
and objectively. Figure 8 shows the stress–strain relationshipobtained using the measured blood pressure pðtÞ and thechange in thickness hðtÞ between the detected LIB and
MAB. The hysteresis loops are similar to those in ourprevious reports.13,14)
Figure 9 shows the means of the measured static elasticity
Es and viscosity  in the cross-sectional scan. The means
were obtained from about 3–5 (cross-sectional scan)
available ultrasonic beams. Horizontal dashed lines show
the means of all of the respective estimates. The means of
the static elasticity Es and viscosity  are 744 kPa and1.55 kPas in the cross-sectional scan. The measured
viscoelasticity was in the same range as that of the carotid
artery reported in the literature (in vitro measurement).20)
Figure 10 shows the transient changes in the means ofstatic elasticity Es and viscosity  averaged by 3–7 available
ultrasonic beams. These parameters were estimated from thestress–strain relationship of the radial artery. The transientchange in static elasticity Es was similar to that in adiﬀerent elastic parameter reported in the literature.11,12)
Additionally, including viscosity, the transient changes inthese parameters were similar to those in our previous
study.14) The minimum static elasticity Es was measured at

Fig. 7. (Color online) In vivo results of the template matching on B-mode
images of a radial artery for 10 min.

7 s after the release of the cuﬀ. The maximum % change instatic elasticity Es was about 71.8% (488 kPa). Moreover,the viscosity , which was evaluated noninvasively by the

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manually detected LIB and MAB. Figure 11 shows thetransient changes in the estimated amplitude [Fig. 11(a)],mean squared error [Fig. 11(b)], and the LIB and MAB
determined by the proposed method and an operator(referring to B-mode images) [Fig. 11(c)]. In Fig. 11(c),the RF signals obtained at the time of the R-wave of theelectrocardiogram are also shown. The LIB positions
detected by template matching were similar to those detectedby the operator. However, after recirculation, signiﬁcantdiﬀerences were observed between the positions detected bytemplate matching and the operator.
Figure 12 shows the transient changes in the static
elasticity Es and viscosity  obtained from the boundaries
detected by the operator. The transient changes in static
elasticity Es and viscosity  were similar to those in Fig. 10.

evaluation of reproducibility at rest
2

4

(a)

6

(b)

(c)

8

(d)

[min]

10

Blood Pressure [mmHg]

130

(a)
(b)
(c)
(d)

17.3

110

14.7

90

12.0

70

9.3

50

0

−2

−4

−6
−8
Strain [%]

−10

−12

−14

Stress [kPa]

0

6.7
−16

Fig. 8. (Color online) Stress–strain relationships obtained from data in

1200

3

1000

2.4

Viscosity [kPa s]

Elasticity [kPa]

Fig. 7.

800
600
400
200
0

1.8
1.2
0.6
0

2 4 6 8 10Time [min]2 4 6 8 10Time [min]Fig. 9. (Color online) Means of static elasticity Es and viscosity  of theradial artery for 10 min at rest. Red dashed lines show the means of all of therespective estimates.

1000

Elasticity [kPa]800

8elasticity

600

6

400

viscosity

200
0

4

Viscosity [kPa s]

10

2

0

60 120 0

30

60

90

120

150

0
180
Time [s]Fig. 10. (Color online) Transient changes in means of estimated static
elasticity Es and viscosity  obtained by boundaries detected using the
proposed template matching method.

proposed method, increased after recirculation. The maximum viscosity  at 114 s after recirculation was about
5.6 kPas, which is about 398% (4.4 kPas) larger than themean at rest.
For the evaluation of the template matching method, thetransient change in viscoelasticity was obtained using the

4. Discussion

In this study, we calculated the MSE between the model andmeasured echoes (not envelope). Figures 6 and 7 show theresults of the boundary detection for several ultrasonicbeams in a B-mode image in the short axis plane of theradial artery. Although Fig. 6(d) shows an example ofultrasonic echoes in which the boundaries of the arterial wall
are diﬃcult to be determined manually, as shown in
Fig. 6(c), the boundary positions that are shown as the LIB
and MAB on the B-mode image are detected appropriately
and objectively by template matching.
Figure 9 shows the means of static elasticity Es andviscosity  measured every 1 min for 10 min at rest. For astable measurement, the ultrasonic beam should be perpendicular to the arterial wall and the intima–media region
should be visualized clearly during the measurement. In
the cross-sectional scan, several ultrasonic beams bisect
at right angles, whereas the number of ultrasonic beams,
which perpendicularly pass through the artery wall,decreases. However, Fig. 9 shows a large ﬂuctuation ineach parameter, even though this method was applied to ahealthy subject. Such a ﬂuctuation should be reduced byfurther investigation.
Figure 10 shows the transient changes in static elasticity
and viscosity. The transient changes in these parameters weresimilar to those in a previous study14) and to those obtainedusing the boundaries detected by an operator (Fig. 12).
This result shows that the proposed method could detectboundaries automatically and objectively. The temporaldecrease in static elasticity Es and the temporal increase inviscosity  were obtained after recirculation. The measured
temporal change in viscoelasticity suggests that this method
is able to measure the temporal change in the mechanicalproperty of the arterial wall due to FMD, although thephysiological meanings of such temporal changes andrecoveries of Es and  should be further investigated.
Figure 11(c) shows the LIB and MAB detected bytemplate matching and an operator (referring to B-mode
images). Just after recirculation, there are signiﬁcantdiﬀerences between the positions detected by templatematching and the operator. Although there are small
ﬂctuations in MSE and the estimated amplitude of echoes
in Figs. 11(a) and 11(b), they are similar just afterrecirculation. However, at the time of 161 s after recirculation, the estimated amplitudes of echoes are similar, but

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(a)amplitude 3
C3
C2

2
1

C1
0

(b)
MSE

1
0.8
0.6
0.4
0.2
0

(c)

4.98 depth
[mm]
5.21

LIB
(template matching)
LIB (manual)

5.44
5.67

MAB (manual)MAB(template matching)

5.91

0

20

40 60 80 100 120[s]
at rest

7

20

32

44

56

67

79

91

after avascularization

103 114 126 138 150 161 172[s]Fig. 11. (Color online) (a) Transient changes in the estimated amplitude, (b) mean squared error, and (c) boundaries determined by the proposed method
and an operator.

Elasticity [kPa]800

elasticity

10

Acknowlegment

8

This work was supported by a Grant-in-Aid for JSPS
Fellows.

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07HF08-7# 2011 The Japan Society of Applied Physics

Fig. 12. (Color online) Transient changes in means of estimated static
elasticity Es and viscosity  obtained using boundaries detected manually.

there is a small diﬀerence between the boundaries detectedby template matching and the operator. Therefore, it isnecessary to reveal the origin of the redundant echo.
As described above, there are many aspects of this study
that should be further investigated. However, the boundaries
were detected objectively and automatically by the proposed
template matching method, and the obtained results showthat the method described in this paper can detect changes inthe viscoelastic properties of the arterial wall due to FMD.
Such a method would be beneﬁcial for the detection of earlystage atherosclerosis.

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5. Conclusions

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In this study, for an accurate and stable measurement oftransient change in viscoelasticity, we applied the proposed
template matching method between the measured RF andadaptive model signals for the boundary detection of theartery. The proposed template matching method showed apotential for the accurate boundary detection and stable
analysis of the transient changes in the mechanical properties of the intima–media region caused by FMD.

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