Japanese Journal of Applied Physics
Vol. 45, No. 5B, 2006, pp. 4718–4721#2006 The Japan Society of Applied Physics

Spatial Distribution Measurement of Heart Wall Vibrations
Generated by Remote Perturbation of Inner Pressure
Hiroshi K ANAI, Hideyuki H ASEGAWA and Kohsuke I MAMURA
Department of Electronic Engineering, Graduate School of Engineering, Tohoku University, Aramaki-aza-Aoba 6-6-05, Sendai 980-8579, Japan(Received November 30, 2005; accepted February 13, 2006; published online May 25, 2006)

It is essential for the diagnosis of heart diseases to noninvasively measure instantaneous myocardial movability and transition
properties during one cardiac cycle. This study proposes a novel method of noninvasively perturbing left ventricle (LV)internal pressure by remotely actuating the brachium artery with sinusoidal vibration for the diagnosis of myocardial
movability. By attaching an actuator to the brachium artery and driving it with a sinusoidal wave of f0 Hz, the internal pressureof the artery is perturbed. The perturbation propagates along the artery to the LV of the heart and the sinusoidal perturbation ofthe LV internal pressure is induced. Using an ultrasound-based phased tracking method, the resultant minute motion of theheart wall can be noninvasively measured. Because the vibration mode of the heart wall depends on actuation frequency, byphantom experiments using a spherical shell made of silicone rubber, to which a silicone rubber tube is connected, thevibration mode was identiﬁed from the measurement of the spatial distribution of the motions by scanning with an ultrasonicbeam. From an in vivo experiment, the principle of remote actuation was conﬁrmed. [DOI: 10.1143/JJAP.45.4718]KEYWORDS: phased tracking method, myocardium vibration, myocardium movability, remote actuation, viscoelasticity, tissue
characterization, internal pressure, vibration mode

1.

Introduction

Though myocardial movability is important in thediagnosis of heart diseases,1–3) a non-invasive method ofestimating the movability has never been developed. The
viscoelasticity of the heart wall is estimated by the ratio of
the applied change in internal pressure to the resultant strain
in the heart wall. However, because large vibration components are contaminated by the heart wall vibration naturallygenerated by a heartbeat, ultrasound-based actuation4–6)
cannot easily be applied under safety guidelines.
When the heart wall is sinusoidally actuated using a largescale low-frequency actuator directly attached to the chest
wall, the bending vibration (mode 2) dominantly occurs in
the heart wall as shown in Fig. 1(c). For mode-2 vibration,
however, there are eight nodes that do not move on the heartwall. Thus, it is necessary to distinguish these nodes fromthe diseased regions to avoid underestimating the movabilityof the myocardial region. For this purpose, it is necessary to
select an actuation frequency f0 so that vibration withmode 0 [as in Fig. 1(a)], where the node is not included, isactuated.
In our laboratory, the change in the internal pressure of thecommon carotid artery (CCA) was generated by remote
cyclic actuation at the brachium artery and the elastic
modulus of the CCA wall was measured by an ultrasoundbased method at multiple frequencies in basic7) and in vivo
experiments8) to non-invasively estimate the viscoelasticityof the arterial wall.9)
In this study, remote actuation is applied so that theinternal pressure of the left ventricle (LV) is successfully
perturbed in in vivo experiments. At the same time, thevelocity signals on the interventricular septum (IVS) aremeasured using the ultrasound-based phased tracking method.10,11) Prior to the in vivo experiments, using a phantom ofa silicone rubber shell similar in size and elasticity to an
actual heart of a typical healthy adult, f0 is predetermined inphantom experiments so that vibration with mode 0 isE-mail address: kanai@ecei.tohoku.ac.jp(a) mode−0(b) mode−1

node

nodenodenodenode(c) mode −2
node

node

node

node

nodenode(d) mode−3

Fig. 1. Each vibration mode with spherical shell.

generated. There is not node in mode-0 vibration. Thismethod has a potential for non-invasively evaluating
myocardial movability and the transition properties of theheart wall during one cardiac cycle.
2.

Principle of In Vivo Actuation

As shown in Fig. 2, by attaching a large-scale lowfrequency actuator to the right brachium artery and driving itwith a sinusoidal wave of f0 Hz (approximately 10 Hz), theinternal pressure of the artery is perturbed and theperturbation propagates along the artery to the LV of the4718

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H. KANAI et al.

function generator

ultrasounddiagnosis
systemsiliconerubber tube

oscilloscope

ultrasonic
probe

ac
tua
tor

rig

ht

br
ac

hium

fixed

work stationx–stagelarge
scaledactuatorcatheter
stage
water
sheet

silicone shelliron board

25 mm 25 mm

ultrasonicprobe

Fig. 2. Photograph of in vivo experiments in which right brachium arteryis actuated with large scale actuator and velocity waves on heart wall aremeasured.

A3.

Basic Experiments Using Silicone Shell

As a model of the heart of a typical healthy adult, we useda spherical shell (outer diameter = 45 mm, thickness ¼7:5 mm) made of silicone rubber (Toshiba YE5623) and set
in a water tank as shown in Fig. 3. A silicone rubber tube
(50 cm long) was connected to the spherical shell. The shelland tube were ﬁlled with water. By remotely actuating thetube using a large-scale actuator with a sinusoidal vibrationof f0 Hz, the internal pressure of the shell was perturbed atseveral mmHg. According to the mechanical test, theelasticity of the silicone rubber was approximately 46 kPa
at 10 Hz, which is similar to the elasticity of the heart wall ofa healthy young subject measured in in vivo experiments(approximately 30 kPa for 10 – 90 Hz).13) The LV is mainly
supported by the aorta. Thus, the boundary condition of thephantom in the experiment is similar to that of the actualheart.
As shown in the lower inset of Fig. 3, for measuring thespatial distribution of the velocity wave on the shell walls,
the ultrasonic probe was set at one of three positions (A, B,and C) using the x-stage. For each position, the ultrasonicC
53 mm

x−stageheart through the aortic valve. Then, the sinusoidal perturbation of the LV internal pressure is generated. Using anultrasound-based phased tracking method,10,11) the resultant
minute motion of the heart wall can be measured. Because
the vibration mode of the heart wall depends on f0 ,myocardial movability cannot be correctly measured atnodes. Thus, the generation of the vibration with mode 0,
which does not include nodes, is essential. For the actual
heart, however, because there are large motion components
due to a heartbeat, it is not easy to accurately detect thespatial distribution of the minute wall motion caused byremote actuation at present. In particular, for the LV
posterior wall, because data at points deep in the skin isrequired, frame rate cannot be increased and accurate
measurement is not easily realized. For the above purpose,
therefore, by the basic experiments described in the next
section, the f0 of vibration with mode 0 is predetermined inphantom experiments from the measurement of the spatial
distribution of heart wall motion by scanning with anultrasonic beam.12,13)B

8 beams
22°

75 mm
siliconeshell
beam 8beam 1

Fig. 3. Upper: Setup for basic experiment, in which internal pressure ofsilicone spherical shell was perturbed and spatial distribution of velocity
waves along shell surface was measured by scanning ultrasonic beams.
Lower: For measuring spatial distribution of velocity wave on shell walls,
ultrasonic probe was set at one of three positions (A, B, C) using x-stage
and ultrasonic beam was scanned sparsely in eight directions at eachposition.

beam was scanned sparsely in eight directions. Thus,
remotely actuated vibrations were measured at 24 points
on the anterior wall and at 24 points on the posterior wall of
the shell. By referring to the actuation signal, the phases inall these measured vibration signals were matched.
Figure 4 shows the waveforms of the velocity signals
simultaneously measured at eight points on the anterior walland at eight points on the posterior wall by setting the probe
at position C. In both walls, there were spatial changes in thephase of the actuated pressure perturbation signal along thewall surface.
At 18 ms in Fig. 4, the amplitudes of each waveform at 48
points along the walls were detected. The results are shownfor f0 ¼ 17 Hz in the upper inset of Fig. 5 and forf0 ¼ 30 Hz in the lower inset of Fig. 5. From the measured
spatial distribution of the vibrations on the shell surface, thevibration mode was identiﬁed. That is, for f0 ¼ 17 Hz, thevibration mode is mode 2, whereas for f0 ¼ 30 Hz, thevibration mode is similar to mode 3.
Figure 6 shows the frequency characteristics of themaximum amplitude of the velocity waveform. For f0 ¼
7 { 8 Hz, the vibration-mode is 0, which indicates thehomogeneous expansion and contraction of the shell wallwithout any nodes as shown in Fig. 1(a). For 12 Hz, thevibration mode is 1, which indicates the parallel displacement of the whole shell, as shown in Fig. 1(b). For f0 ¼
16 {17 Hz, the vibration mode is 2. From this distribution,
the positions of the nodes were identiﬁed.

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2520151050−5−10−15−20− 250H. KANAI et al.

velocity [mm/s]

45

anteriori wall

beam 1mode 340
mode 2velocity [mm/s]

35

beam 8206025
mode 020
151080100time [ms]mode 150posterior wall

velocity [mm/s]

2520 beam 1
151050−5−10−15− 20 beam 8−25020403005101520
frequency [Hz]2530Fig. 6. Maximum amplitude of velocity waveform measured for eachapplied f0 .

4060

80
100
time [ms]interventricularseptum

Fig. 4. Waveforms of velocity signals simultaneously measured at eight
points on anterior wall and eight points on posterior wall by setting probe
at position C.

right ventricle
measured points

left ventricle
left atrium+10 mm/s
beam 1beam 8beam 8–10 mm/s

Abeam 1

Fig. 7. Cross-sectional image of heart and employed beam direction.

beam 8beam 1

BC+10 mm/s
–10 mm/s

ABC

Fig. 5. Amplitude of each vibration waveform at 48 points, detected attime of 18 ms in Fig. 4. Upper: f0 ¼ 17 Hz. Lower: f0 ¼ 30 Hz.

4.

In Vivo Experiments

By applying the external remote actuator shown in Fig. 2
to the brachial artery of a healthy male subject (32 year old),
LV internal pressure was successfully perturbed with mode0 vibration. The f0 was approximately 9 Hz, which corresponds to mode 0 in Fig. 4. There was suﬃcient reproducibility. The motion of the heart wall as measured from thechest wall using the ultrasound-based phased tracking

method.10,11) Figure 7 shows a cross-sectional image of theheart and the employed beam direction.
Figures 8(e) and 8(f) show typical waveforms of thevelocity signals during the actuation fvact ðt; iÞg (i = LV, RV)
and the cessation of the actuation fvces ðt; iÞg (i = LV, RV)
on both sides, that is, the right-ventricle (RV) side and theLV side of the IVS during one cardiac cycle. Figure 8(g)shows typical waveforms of the change in the thickness ofthe IVS during the actuation fhact ðtÞg and the cessation of theactuation fhces ðtÞg, which are obtained by integrating thediﬀerences between these velocity signals on both sides ofthe IVS. The red lines and blue lines in Figs. 8(c)–8(g) show
the vibrations or change in thickness measured when thebrachium artery was actuated and when it was not actuated,respectively.
Each of the velocity signals was averaged during theactuation fvact ðt; iÞg (i = LV, RV) and the cessation of theactuation fvces ðt; iÞg (i = LV, RV), and the averaged differences, fvðt; iÞ ¼ vact ðt; iÞ  vces ðt; iÞg (i = LV, RV), between these averaged velocity signals, fvact ðt; iÞg (i = LV,RV) and fvces ðt; iÞg (i = LV, RV), are shown by solid lines
for the RV and LV sides in Figs. 9(c) and 9(d), respectively.
The broken lines in Figs. 9(c) and 9(d) show the actuation
signal that drives the actuator. By comparing these signals,
the diﬀerence (solid lines) almost corresponds to theactuation signal (broken lines) particularly during systole(marked at three times by arrows) as shown in Figs. 9(c) and9(d).

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right ventricle

(a)
(b)

interventricular septum

systole5 mm/sleft ventricle(c)0depth 5.1 mm(a)

5 mm/s(d)0right ventricle
interventricular septum

0.2 mm(e)0left ventricle

−200

(b)

0(c)
(d)

(e)
50 mm/s(f)

04 mm−2000200
400
600
800
time from R−wave [ms]600

During systole, the aortic valve is open so the perturbation of
the internal pressure due to actuation at the brachial artery
passes through the aortic valve and propagates into the LV.

(g)

0200
400time from R−wave [ms]Fig. 9. Diﬀerence between waveforms respectively averaged during
actuation and cessation in Fig. 8. The broken lines: the averaged
actuation signals which electrically drives the actuator. (a) ECG withactuation. (b) ECG without actuation. (c) Averaged diﬀerence (solid line)vðt; RVÞ between velocity signals vact ðt; RVÞ and vces ðt; RVÞ at RV sideof IVS respectively averaged during actuation and cessation. (d)Averaged diﬀerence (solid line) vðt; LVÞ between velocity signals
vact ðt; LVÞ and vces ðt; LVÞ at LV side of IVS respectively averaged during
actuation and cessation. (e) Averaged diﬀerence (solid line) hðtÞ
between averaged changes in thickness, hact ðtÞ and hces ðtÞ, of IVS
averaged during actuation and cessation.

50 mm/s

001000Fig. 8. Waveforms during four cardiac cycles in in vivo experiments.
(a) M-mode and tracking lines (red) with actuation. The f0 is 9 Hz.
(b) M-mode and tracking lines (red) without actuation. (c) ECG.
(d) Actuation signals. (e) Velocity signals fvact ðt; RVÞg and fvces ðt; RVÞgat RV side of IVS. (f) Velocity signals fvact ðt; LVÞg and fvces ðt; LVÞg atLV side of IVS. (g) Change in thickness fhact ðtÞg and fhces ðtÞg of IVS. For(c)–(g), the red lines correspond to the experiments with actuation andblue lines correspond to the experiments with cessation of the actuation.

In a similar way, for the change in thickness, the averaged
diﬀerence hðtÞ between fhact ðtÞg and fhces ðtÞg is shown by asolid line in Fig. 9(e). By comparing the solid line with thebroken line in Fig. 9(e), the change in thickness hðtÞ (solidline) has an almost opposite phase to that of the actuation
signal (broken line) at the times marked by arrows.
For f0 of 9 Hz, the minute velocity components of 5 mm/s
due to the perturbation in Figs. 9(c) and 9(d) of the internalpressure were superposed on the spontaneous myocardial
motion of approximately 50 mm/s in Figs. 8(e) and 8(f).
Also, for the change in thickness of the IVS, minute changes
in thickness of approximately 200 mm were superposed on
spontaneous changes in thickness of approximately 4 mm inFig. 8(g). The ratio of the change in thickness caused byremote actuation to the spontaneous change in thickness isapproximately 1 : 20 and the pulse pressure is approximately
40 mmHg. Thus, the change in pressure in the LV caused byremote actuation is approximately 2 mmHg. A similarchange in pressure has been measured for the CCA using
the same remote actuation system in in vivo experiments.8)

5.

Conclusions

In this study, we propose a novel method of noninvasively perturbing the LV internal pressure by remotelyactuating the brachium artery with sinusoidal vibration. By
measuring the spatial distribution of phantom motion using
an ultrasound-base method, the vibration mode is identiﬁed,
which has the potential for non-invasively evaluating themyocardial movability and the transition properties of theheart during the cardiac cycle.
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