REGULAR PAPER

Japanese Journal of Applied Physics 53, 07KF08 (2014)
http://dx.doi.org/10.7567/JJAP.53.07KF08

Echo speckle imaging of blood particles with high-frame-rate echocardiography
Hiroki Takahashi1*, Hideyuki Hasegawa1,2, and Hiroshi Kanai1,2
1Graduate School of Biomedical Engineering, Tohoku University, Sendai 980-8579, Japan
Graduate School of Engineering, Tohoku University, Sendai 980-8579, Japan
E-mail: htaka@us.ecei.tohoku.ac.jp
2

Received November 26, 2013; revised February 21, 2014; accepted March 5, 2014; published online June 13, 2014
Cardiac blood ﬂow patterns such as the vortex ﬂow pattern inside the left ventricle have been studied to provide new information for the diagnosisof the pumping function of the human heart. Complex blood ﬂow is visualized by imaging echo speckles of blood particles because the speckle-liketexture translates to the motion of blood particles. We proposed an imaging method for echo speckles of blood particles using high-frame-rate
ultrasound for the visualization of the intracardiac blood ﬂow direction. High-frame-rate ultrasound is useful for continuously observing the fast
motion of echoes from blood particles in the heart. In the present study, weighting by coherence and compounding the magnitudes of echo signals
in different transmissions were introduced to visualize weak echoes from blood particles. The feasibility of the visualization of cardiac blood ﬂow
using high-frame-rate ultrasound was demonstrated by basic and in vivo experiments. © 2014 The Japan Society of Applied Physics

1.

Introduction

Echocardiography is a predominant imaging modality for theevaluation of the cardiac function because it could provide across-sectional image of the heart noninvasively in real time.
Recently, it has been shown that the ultrasonic measurementof myocardial contraction and its propagation is useful for theevaluation of the myocardial function.1–3) On the other hand,the evaluation of blood ﬂow structures in echocardiography
is also useful for the diagnosis of the pumping function of
the human heart.4,5) As a typical method, color Doppler ﬂow
imaging (CDI) has been widely used to obtain information
on intracardiac blood ﬂow in clinical practice. CDI, however,does not show the direction of blood ﬂow because it provides
only the velocity component along an ultrasonic beam.
Recently, cardiac blood ﬂow patterns such like the vortex
ﬂow pattern in left ventricular cavity have been studied toprovide new information for diagnosis by contrast echocardiography, which is called echocardiographic particle image
velocimetry (E-PIV).6–8) E-PIV gives the distribution ofvelocity vectors of blood ﬂow, which is calculated on thebasis of the motion of ultrasonic echoes from contrast agents.
E-PIV, however, creates physical and mental burden inpatients owing to the intravenous injection of a contrast
agent.
An estimator of the velocity vector of blood ﬂow has beendeveloped by several researchers. Tortoli and coworkers9,10)
proposed a method of determining the angle of blood ﬂow,
where a beam is used for estimating the ﬂow direction byidentifying the angle that shows symmetrical Doppler spectrawith respect to the zero frequency. Another beam is used toobtain the angle-corrected Doppler spectra. Jensen and
Nikolov11) developed a method of ﬂow vector estimation
using a synthetic aperture technique. In their method, echo
signals are created along the manually determined ﬂow
direction and the ﬂow velocity magnitude is estimated byapplying a cross-correlation technique to echoes obtained bythose beams. These methods are intended for the measurement of blood ﬂow inside a vessel where blood particles
move on a preassigned streamline along the blood vessel.
However, the blood ﬂow inside the cavity of the heart
exhibits complex ﬂow patterns, such as a vortex pattern, in
contrast to the vascular ﬂow. Furthermore, blood particles
ﬂow at a velocity of over 1 m/s in the human heart. Hence, amethod of scanning a broad region that covers the ventricular
cavity in a short time, i.e., high-frame-rate imaging, isrequired to visualize the complex and fast ﬂux in the humanheart. High-frame-rate ultrasound with parallel beamforming,12,13) therefore, would be of value for the visualizationof a rapid and complex blood ﬂow. Some authors havesuggested methods of visualizing the direction of blood ﬂow
in a blood vessel by analyzing the motion of echoes fromblood particles using parallel beamforming with plane wave
transmission.14–16) Moreover, the intracardiac vector ﬂow ofneonates and children was visualized by high-frame-rate
ultrasound with plane wave transmission using a linear arrayprobe.17) However, a phased array probe is normally used
for the sector scanning of an ultrasound beam in cardiac
ultrasound. The use of plane waves, hence, limits the scan
speed in cardiac ultrasound because the width of insoniﬁed
plane waves does not increase with the range distance fromthe probe in a near ﬁeld, whereas the lateral interval of scanlines increases with the range distance, and the width of planewaves decreases with an increase in the steering angle.
Hasegawa and Kanai developed an alternative method ofhigh-frame-rate echocardiography using parallel beamforming with spherically diverging waves emitted with alltransducer elements.18,19) These spherical diverging waves
used in this method do not limit the illuminated region by
the size of the aperture and enable the acquisition of echoeswith higher signal-to-noise ratios (SNRs) than the synthetic
aperture ultrasound,11,20) which uses one or small number of
elements.
We proposed a method for the visualization of theintracardiac blood ﬂow direction using parallel beamformingwith the transmission of spherically diverging waves. Coded
excitation was used to enhance weak echoes from blood
particles. Echoes from blood particles were extracted using
high-pass ﬁltering for clutter suppression. In addition tothese conventional techniques, two unique approaches were
proposed to enhance echoes from blood particles in the
proposed method. To suppress undesired noise components
in the ﬁltered echo signals, high-pass ﬁltered echo signals
were weighted using a coherence function in the ﬁrst
approach. Compounding was applied between transmissions
to envelope signals in the second approach because echo
signals from moving blood particles obtained in different
transmissions cancel each other when compounding is

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H. Takahashi et al.
distanse

Jpn. J. Appl. Phys. 53, 07KF08 (2014)rf

sector probeelementθ
focal pointtransmit waveform

transmit

receive
400 μmFig. 1. (Color online) Illustration of transmission of spherically diverging
beam and parallel receive beamforming using phased-array probe.

receiveinsonify

applied to RF signals. The accuracy of our proposed method
was evaluated by a basic experiment using blood-mimicking
ﬂuid ﬂowing inside a tube. Furthermore, the feasibility ofthe proposed method was demonstrated by an in vivo
experiment.
2.

Principles

2.1 Parallel receive beamforming with sphericallydiverging wave for high frame rate echocardiography

In this study, parallel receive beamforming with sphericallydiverging beams (PBF-DB), which is shown in Fig. 1, wasused. Using PBF with broad transmit beams, echo signals
are measured at a high frame rate without sacriﬁcing thedensity of scan lines. A spherically diverging wave, which
can illuminate a wider region, was transmitted to solve theproblem of the limited width of a plane wave. An ultrasonicimage of a blood pool, which is ﬁlled with blood particles(scatterers), has a speckle-like texture21,22) because theconstructive and destructive interferences among echoesfrom many scatterers much smaller than the wavelength oftransmitted ultrasound occur in the focal point, as shown in
Fig. 2. Thus far, many researchers have examined methods
utilizing an echo speckle, which contains information on
tissue structures for blood ﬂow vector estimation,23) thequantitative evaluation of liver ﬁbrosis,22,24) and shear wave
velocity measurement for tissue elasticity imaging.25–27) An
echo speckle, which arises from blood echoes, translates
between successive frames, i.e., the spatial feature of anecho speckle is preserved, when blood particles move in thesame direction without changing their relative distance. The
velocity vector of a slow blood ﬂow in the left atrium has
been calculated using the dynamic texture produced bystrong backscattering.28) However, the speckle texture
frequently transmutes during a short time because blood
particles in the left ventricular cavity ﬂow at a velocity of
up to 1 m/s. PBF-DB is helpful for preserving the spatial
feature of echo speckles during a frame interval because theframe rate is further increased by imaging a wider region by
one transmission as compared with parallel beamformingwith plane wave transmission. In Fig. 1, rf and ª are thedistance from a virtual point source behind the array to thecenter of the transducer array and the steering angle,
respectively. A spherically diverging beam is formed by
activating each transducer element with a time delay
calculated using rf and ª.18,19)

scatter
8 μm

blood particles

Fig. 2. Illustration of echo waveform from region composed of many
blood particles, i.e., many scatterers much smaller than the wavelength at an
ultrasonic frequency of 3.75 MHz. Echo signals from blood particles areindistinguishable from each other owing to the interference of the scatteredsignals.

2.2 Coded excitation for enhancement of echoes fromblood particles

The amplitude of a scattered signal depends on the size ofthe scatterer, the reﬂection coefﬁcient of the surface of thescatterer, and the frequency of the transmitted ultrasound.29)
Most of the blood particles are red blood cells, which are verysmall ultrasonic scatterers whose diameter is almost 8 µm.
The reﬂection coefﬁcient between the blood particle andblood plasma is also small.30) It is well known that the powerof echoes from scatterers like red blood cells is proportional
to the fourth power of frequency. It is, however, necessary totransmit low-frequency ultrasound waves (from 2 to 5 MHz)in cardiac ultrasound to get echoes with better SNRs from theprofound heart by avoiding signiﬁcant frequency-dependent
attenuation. Hence, the enhancement of echo signals fromblood particles is essential for the visualization of echo
speckles of blood particles in cardiac ultrasound. SNRs ofecho signals from blood particles can be increased bytemporal (frame) averaging. Temporal averaging, however,degrades the spatial features of echoes from blood particles
because echo speckles of blood particles are blurred owing to
the motion of blood particles during temporal averaging. On
the other hand, the SNRs of echo signals could be improvedby increasing the transmit pulse amplitude, but the transmit
pulse amplitude is limited by safety regulations. Bi-phasecoding, such as Barker and Golay coding, which improves
the SNRs of signals by the compression of longer pulses
without safety violations, could be realized at a low cost.31)
Since Golay coding requires two transmissions, the compression gain would be further degraded by the tissue motion
between two transmissions. Hence, in the present study,
coded excitation and pulse compression with a Barker
code,32) which is a single-transmit code, were used to

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H. Takahashi et al.

increase the SNRs of weak echoes from blood particles. The
excitation signal was coded with a 13-bit Barker code.
Moreover, the echo signal received at each element was
individually decoded before receive beamforming.
High-pass ﬁltering for clutter suppression
Echoes from stationary and slowly moving tissues such as the
ribcage and myocardium caused by the sidelobe, which arecalled clutter signals, contaminate those from blood particles
in the cardiac lumen. Suppressing the clutter noise is essential
for extracting echoes from blood particles because theamplitudes of clutter echoes are usually much larger thanthose of echoes from blood particles. On the other hand, themotion velocity of the myocardium is different from that ofblood particles. The maximal velocity of the myocardium isalmost 0.1 m/s, whereas that of blood ﬂow in the humanheart reaches to 1 m/s. Clutter signals, hence, have a lowerDoppler-frequency component than echo signals from blood
particles, which corresponds to a small change in the phase ofan echo signal in the direction of the frame. In the present
study, clutter signals were suppressed by high-pass ﬁlteringin the direction of the frame. The ﬁlters’ impulse response
was the sinc function multiplied by the Kaiser window, which
is linear time-invariant. Myocardial walls normally movewith a velocity of up to 0.1 m/s,33) which corresponds to a
Doppler frequency of 500 Hz in the case of a transmission
frequency of 3.75 MHz. To suppress clutter signals generated
by sidelobes, the cutoff frequency of a clutter ﬁlter was set
to 500 Hz.
2.3

2.4

Weighting by coherence for noise suppression
The undesirable noise component, which is generated byelectronic devices, still remains in the high-Doppler-frequency range although a clutter ﬁlter is used. The suppressionof the remaining noise is efﬁcient for the visualization ofechoes from blood particles because the noise component
contaminates weak echoes from blood particles.
The temporal changes in the phase of the signal component(corresponding to echoes from an object) would be spatiallyconstant during a pulse duration, i.e., coherent. The temporalchanges in the phase of the signal component at a certain
depth is also assumed to be constant for a short period (small
number of frames). On the other hand, the temporal changes
in the phase of the noise component are random, spatiallyand temporally. Therefore, in the present study, the random
noise remaining in the high-Doppler-frequency range wassuppressed by weighting using the ratio of the coherent
component to the total signal. The coherent component was
obtained by minimizing the noise component using the leastsquare method applied to quadrature demodulated signals.
Let us deﬁne the complex demodulated signal at the frame
number i and the sampled point number n, which corresponds
to time t = nT (T: sampling interval) as xi(n). The normalized
mean square error (corresponding to the ratio of the noise
component) ¡ among complex RF signals fxi(n)g for a few
sampled points Nd and frames Nf is expressed as follows:¼

2

Ek ½Em ½jxiþm ðn þ kÞ  z  xiþm1 ðn þ kÞj ;
Ek ½Em ½jxiþm ðn þ kÞj2 ð1Þwhere z, Ek[0], and Em[0] denote the transfer function fromfxi¹1(n)g to fxi(n)g and the averaging procedures with respect

to the depth and frame, respectively. The averaging numbersof Ek[0] and Em[0] are deﬁned by Nf and Nd, respectively. The
estimated z^ , which minimizes ¡, can be estimated by the leastsquare method using Eq. (1). By substituting z^ into Eq. (1),the minimum mean square error ¡min is obtained as follows:
jEk ½Em ½xiþm1 ðn þ kÞxiþm ðn þ kÞj2:
Ek ½Em ½jxiþm1 ðn þ kÞj2 Ek ½Em ½jxiþm ðn þ kÞj2 ð2ÞThe obtained ¡min corresponds to the normalized residual
noise component in the signal fxi(n)g, which cannot be
predicted by fxi¹1(n)g. Hence, the instantaneous intensity of
the coherent component at the n-th depth point in the i-thframe is obtained as (1 ¹ ¡min), and the estimate S^ i ðnÞ ofthe amplitude of echo from blood particles is obtained byweighting the
instantaneous
signal amplitude «xi(n)« with thepﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃ
ﬃ
coherency 1  min as follows:
pﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃ
S^ i ðnÞ ¼ jxi ðnÞj ð1  min Þ:
ð3Þmin ¼ 1 In the present study, the lengths for averaging in the frameand depth directions, Nf and Nd, were set to 6 and 8(corresponding to about 2 ms and 0.5 µs), respectively. The
random noise contained in high-pass-ﬁltered echo signals
would be suppressed by weighting the echo amplitude withcoherency.
2.5 Compounding of speckle images obtained bydifferent transmissions

In PBF with a wide transmit beam for high-frame-rateechocardiography, discontinuities of echo intensities occur inan ultrasound image because the transmit–receive directivity
signiﬁcantly changes owing to the large interval between
neighboring transmit beams. To solve this problem, compounding echo signals obtained by different transmissions iseffective.18) The compounding procedure, however, would
not ﬁt the visualization of echoes from blood particles
because the phase of the echo signals from fast moving
objects, such as blood particles, differs signiﬁcantly amongemissions and coherent compounding would be difﬁcult.11)
Incoherent compounding weakens echoes from blood
particles. In the present study, the envelope of high-passﬁltered RF signals in each transmission, which removes
phase information, was compounded (summed) to avoid suchcompensation. Figure 3 shows a schematic diagram of the
proposed method. Let us deﬁne this technique for blood echo
speckle visualization “magnitude compounding”. At ﬁrst,
clutter ﬁltering and weighting by coherence were applied
to beamformed RF signals obtained from each transmission.
Then, from the ﬁltered RF signals, the signal amplitudes[corresponding to S^ i ðnÞ in Eq. (3)] were extracted. Finally,
the amplitudes fS^ i ðnÞg obtained for multiple transmissions
were compounded. The effect of magnitude compounding
was veriﬁed in the subsequent section.
3.

Materials and methods

A blood-mimicking ﬂuid (Shelly Medical Imaging Technologies US model) that was made to ﬂow through a cylindricaltube was measured by ultrasound. The internal and external
diameters of the tube made of urethane rubber were 8 and
12 mm, respectively. The tube contained 5% carbon powder(by weight) to obtain sufﬁcient scattering from the wall for

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Jpn. J. Appl. Phys. 53, 07KF08 (2014)
RF signals
in 2nd transmission

5 mm

RF signals
in 3rd transmission

wall of tube

individual decoder
receive beamformer

0.4
0.2

flow axis

clutter filtering

compound for B−mode image

0.6
velocity [m/s]RF signals
in 1st transmission

H. Takahashi et al.

0envelope detection

Fig. 4. (Color online) Echo speckles overlaid on B-mode image andvelocity vector distribution along direction perpendicular to ﬂow axis for
estimation of ﬂow rate of blood-mimicking ﬂuid ﬂowing inside tube.

weighting by coherencecompound for speckle image of blood particle
10 mmTable I.

image for tissue

overlaid image

image for blood flow

Fig. 3. (Color online) Schematic of procedure of proposed method ofimaging of movement of echoes from blood particles.

the identiﬁcation of the walls, between which the bloodmimicking ﬂuid ﬂowed. The blood-mimicking ﬂuid wasmade to steadily ﬂow using a screw pump (Heishin
2NL10PU), which could control the ﬂow rate in proportion
to the rotation rate of the screw. The Reynolds number
reached up to 1070 under the present experimental conditions. To avoid to the generation of tiny air bubbles in theﬂuid ﬂowing inside the tube, the blood-mimicking ﬂuid was
deaerated using a vacuum pump.
The tube and human heart were measured by a modiﬁed
diagnostic ultrasound system (Hitachi-Aloka Medical Alpha10) with a 3.75 MHz phased array probe. This system was
modiﬁed to acquire individual RF echo signals, which were
sampled at 15 MHz, received by 96 elements. The sphericallydiverging wave was transmitted in three directions andsteered at ¹10, 0, and +10°. PBF and the proposed process
for the visualization of blood ﬂow were applied to off-line
processing.
The velocity vector of the ﬂow was estimated by applying
the speckle tracking technique23) to the distributions of echo
amplitudes in two frames, which was obtained using Eq. (3).
The velocity vector was obtained by searching for the
maximum of the two-dimensional cross correlation function
between the echo amplitudes inside kernels in two consecutive frames. The kernel sizes in the lateral and axial
directions in the speckle tracking technique were set to 9° and4.6 mm, respectively. Note that the cross correlation map was
up-sampled34) by factors of 120 and 10 in the lateral and axial
directions, respectively, before calculating the cross correlation function.
4.

Basic experimental result

The distance between the outer surface of the tube anteriorwall and the probe surface was set at 5 cm. The angle
between the probe surface and the tube was maintained at
15°. The tube’s longitudinal direction (corresponding to theﬂow axis) was manually identiﬁed in the B-mode image. The
velocity vectors of the ﬂuid inside the tube were estimatedCross-sectional areas of lumen of tube.

Rotation rate of screw(min¹1)

Cross-sectional area
(cm2)

160

0.56

260
360

0.53
0.44

460

0.38

at 0.5 mm intervals along the manually assigned line by
applying the speckle tracking method to echo speckles fromblood particles obtained by the proposed method. Figure 4
shows a B-mode image of the tube obtained by PBF-DB
overlaid with blood echo speckles and the estimated velocity
vectors. The components of the estimated ﬂow vectors along
the direction of the ﬂow axis were averaged as follows:
Nj1
AX
Q¼
vj  nf ;ð4Þ
Nj j¼0where A, vj, and nf denote the cross-sectional area where theblood-mimicking ﬂuid ﬂowed, the estimated velocity vectorof the j-th spatial point, and the unit vector parallel to theﬂow axis, respectively. The cross-sectional area of the lumenof the tube, A, was changed by the suction pressure of thepump depending on the rotation rate of the screw. Table I
shows the cross-sectional area of the lumen of the tube
measured in the short-axis view obtained with Alpha 10 bymanually segmenting B-mode images to accurately estimate
the ﬂow rate.
The ﬂow rate was estimated by a ﬂow meter (Keyence FDSF1) and our proposed method, while the rotation rate of thescrew pump was varied as 160, 260, 360, and 460/min. The
performance curve of the pump between the rotation rate of
the screw and the ﬂow rate was estimated on the basis ofrelationship between the rotation rate and the ﬂow rate up to
1000 mL/min because the accuracy of the ﬂow meter was notguaranteed when the ﬂow rate was over 1000 mL/min of the
ﬂow rate. Figure 5 shows the ﬂow rate plotted as a functionof the rotation rate of the screw and the regression line
estimated by the least square technique, which is given as
y = 3.4x ¹ 9.2 mL/min [y and x correspond to the ﬂow rate
(mL/min) and rotation rate of screw (/min), respectively].
Figure 6 shows the ﬂow rate estimated by our proposed
method with magnitude compounding (red crosses) and our
proposed method with compounding of RF signals (green
crosses). In Fig. 6, plots and vertical bars denote the means

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H. Takahashi et al.
80

1600

60

1200
800
400
0

40
30

10

0

100 200 300 400
rotation rate of screw [/min]

500

0
100

200

300

400

500

rotation rate of screw [/min]Fig. 7. (Color online) Error of direction of velocity vector estimated byproposed method with RF signal compounding and proposed method withmagnitude compounding.

estimate using flow meter

2.5

magnitude compounding
RF compounding2

1.2

magnitude compounding
RF compounding

1

estimateusing flow meter

flow rate [L/min]flow rate [L/min]

50

20

Fig. 5. (Color online) Flow rates measured at different rotation rates ofscrew and regression line estimated by least square technique. The regression
line is given as y = 3.4x ¹ 9.2 [mL] (y and x correspond to the ﬂow rate
[mL/min] and the rotation rate of the screw [/min], respectively).

1.5

magnitude compounding
RF compounding70

regression line
error [%]

flow rate [mL/min]

2000

1
0.5
0

0.8
0.6
0.4
0.2
0
−0.2

−0.5
100

200

300

400

−0.4

500

rotation rate of screw [/min]Fig. 6. (Color online) Flow rate of blood-mimicking ﬂuid obtained using
ﬂow meter, proposed method with RF signal compounding, and proposed
method with magnitude compounding.

and standard deviations of the ﬂow rate in the temporal
direction for 50 frames. As shown in Fig. 6, the difference
between the means obtained using the proposed methodand ﬂow meter was reduced by magnitude compounding.
From these results, the cancellation of RF signals caused by
the movement of particles degraded the accuracy of theestimation of the velocity vector, and such degradation was
suppressed by the proposed method. The proposed method,
nevertheless, tended to underestimate the ﬂow rate.
The mean error edir of the direction of the estimated ﬂow
vector was also evaluated using the following equations:
qﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃﬃ
Ni
Ej ½ðargðnf Þ  argðvi;j ÞÞ2 
X
1
;
ð5Þ
edir ¼
Ni i¼0
where Ej and arg(0) represent the spatial averaging procedureswith respect to the points of interest assigned along thedirection perpendicular to the ﬂow axis and the argument ofa vector. Figure 7 shows the mean error of the direction of
the ﬂow vector, edir, estimated by the method proposed in thisstudy plotted as a function of the rotation rate of the screw
pump. As shown in Fig. 7, the use of the magnitude
compound could signiﬁcantly reduce the error in theestimated ﬂow direction. Figure 8 shows the ﬂow rate at a
rotation rate of the screw of the 260/min, estimated at angles
between the probe surface and the tube of 25, 15, and 5. The
movement of echo speckles of blood particles could not be

0

5

10
15
20
angle [deg]25

30

Fig. 8. (Color online) Means and standard deviations of ﬂow rateestimated by proposed method at three different angles between the surface
of probe and the tube: about 25, 15, and 5°.

clearly seen at angles of over 25° because the transmitted
ultrasound could not penetrate the tube well owing to thesmall insoniﬁcation angle. The standard deviation of theestimated ﬂow rate was high, but the means were not very
different among the angles of 25, 15, and 5. From theseresults, the accuracy of the ﬂow direction estimated by the
proposed method was improved using magnitude compounding. In this basic experiment, the SNRs of echoes fromparticles mimicking blood particles were much higher thanthose from blood particles in the human heart because theultrasonic attenuation of water was very low. A higher
acoustic output, therefore, would be desired for better
visualization of echo speckles of blood particles in in vivotransthoracic measurement. Furthermore, the velocity anddirection of blood ﬂow rapidly vary in the human heart.
However, the blood-mimicking ﬂuid steadily ﬂowed in thisbasic experiment. Hence, the time length (number of frames)
for clutter ﬁltering, which was set to almost 13 ms in thisstudy, should be as short as possible to suppress contamination by RF signals in other frames, which exhibit a blood
velocity signiﬁcantly different from that in the target frame.
5.

In vivo experimental result

As shown in Fig. 9(a), ultrasonic RF echo signals from theleft ventricle were measured in a three-chamber view of a 26year-old healthy male at a very high frame rate of 2008 Hz.
The healthy subject gave informed consent in the study.

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H. Takahashi et al.

(a)ultrasonic probe

left ventricular cavity

10 mm

septumapical sidebasal side

0.6

aortic valve

(b)

(b)

10 mm

velocity [m/s]posterior wall

0.4
0.2
0

(c)
Fig. 10. (Color online) (a) B-mode image (shown by gray-scale) overlaid
with speckle image of blood particles (shown by hot-scale) at frames in(a) systole and (b) diastole with ﬂow velocity vectors estimated by the
speckle tracking technique.

directed to the aortic valve showed that blood in the left
ventricular cavity was ejected to the aorta in systole, as
shown in Fig. 10(a). On the other hand, as shown in
Fig. 10(b), speckles of blood particles move to the apex sidefrom the basal side in diastole, which corresponds to blood
ﬂowing into the cavity.

(d)

6.

Fig. 9. (Color online) (a) Illustration of measurement region in threechamber view and (b) B-mode image obtained by PBF-DB in a certain framein systole. Speckle images of blood particles (c) without weighting and(d) with weighting by coherence. White dashed lines show the roughly
estimated boundary of the left ventricular cavity.

Figure 9(b) shows the B-mode image obtained using PBFDB in a frame in systole. Figures 9(c) and 9(d) show only
speckle images of blood particles (shown on a hot scale) in
the same frame obtained using the proposed method without
and with weighting by coherence, respectively. The intensity
was individually normalized by the maximums in Figs. 9(c)and 9(d) in that frame. As can be seen in Figs. 9(c) and 9(d),speckle-like patterns of echoes from blood particles werevisualized inside the left ventricular cavity by suppressing
echoes from the myocardium. Moreover, the noise components, which were seen like small and randomly distributed
particles, were suppressed by weighting with coherence, as
can be seen in the region surrounded by the white dashed line
in Figs. 9(c) and 9(d).
Figures 10(a) and 10(b) show ﬂow velocity vectors
overlaid on the B-mode images in frames in systole anddiastole, respectively. The motion of speckle-like patterns

Discussion

The accuracy of the estimation of the velocity vector was
improved in terms of both ﬂow rate (corresponding to ﬂow
velocity) and direction by magnitude compounding, asshown in Figs. 6 and 7, because SNRs of echoes from fastmoving blood particles were improved by avoiding thecompensation caused when RF signals were compoundedincoherently. The mean ﬂow rates estimated using the
proposed method, nevertheless, tended to be lower thanthose estimated using the ﬂow meter, as shown in Fig. 6.
Moreover, the diameter of the tube inhomogeneously
changed in a circumferential direction with an increase in
ﬂow rate (particularly, at a rotation rate of over 360/min)because the elasticity of the wall of the custom-made tube
was inhomogeneous. Therefore, the ﬂow velocity estimatedalong the assigned line was considered to be different fromthe mean ﬂow velocity averaged in the cross section of the
tube. Moreover, in the experimental situation, turbulent ﬂow
was generated because the Poiseuille ﬂow was not fully
developed owing to an insufﬁcient length of the inlet to thetube. On the other hand, the kernel size in the speckle
tracking technique used in the present study, was 7.9 ©
4.6 mm2 at a depth of 5 cm, which was close to the innerdiameter of the tube of 8 mm. The kernel, therefore, could
not follow fast-moving particles because the movements ofparticles were not homogeneous owing to the ﬂow-rate
gradient of the Poiseuille ﬂow and small regional turbulence.
On the other hand, under the in vivo condition, it is also

07KF08-6© 2014 The Japan Society of Applied Physics

Jpn. J. Appl. Phys. 53, 07KF08 (2014)

H. Takahashi et al.

difﬁcult to estimate the velocity vectors of complex blood
ﬂow using a large kernel when blood particles in the kernel
move inhomogeneously. A velocity vector estimator with ahigh spatial resolution is desired to visualize the vector
distributions of complex blood ﬂow in the cardiac cavity. To
solve a such problem, the spatial resolution of the velocity
vector estimator should be improved by considering the useof the optical ﬂow method.35,36) Furthermore, the directional
error decreased with the ﬂow rate, as shown in Fig. 7. The
clutter ﬁlter used in the present study could suppress thesignal components from objects that move with a velocity
less than 0.1 m/s along a scan line. The average ﬂow velocity
along the ﬂow axis was estimated to be 0.41 m/s from theﬂow rate at a rotation rate of 360/min and that along thescan line was also estimated to be 0.11 m/s (over 0.1 m/s
corresponding the cutoff velocity) from the angle between thearray surface and the tube. The directional error, therefore,
would have been reduced because the SNRs of ﬁltered signal
slightly increased with a ﬂow rate under the experimental
conditions used.
7.

Conclusions

In this study, high-frame-rate echocardiography with spherically diverging waves was used to continuously visualize
echo speckles of blood particles. Coded excitation using
the 13-bit Barker code and clutter ﬁltering were also used
to increase the SNR of weak echoes from blood particles.
Moreover, the magnitude of the ﬁltered RF signal was
weighted by coherence to suppress the random noise
remaining after clutter ﬁltering. In the proposed method,
magnitude compounding among transmissions was employed
to obtain echo speckles with higher SNRs because echo
signals from moving blood particles cancel each other when
RF signals were compounded. The ﬂow velocity vector
was estimated by applying the speckle tracking technique tothe visualized blood echo speckle. From the basic experimental result, the accuracy of the estimation of the velocity
vector was improved by magnitude compounding. B-mode
images overlaid with echo speckles of blood particles wereobtained in the three-chamber view through in vivo measurement of a 26-year-old healthy male. Furthermore, the velocity
vectors of blood ﬂow could be visualized inside the left
ventricular cavity by estimating the motion of echoes fromblood particles between successive frames. The in vivo
experimental results showed the feasibility of the noninvasive
imaging of cardiac blood ﬂow patterns using our proposed
method.
Acknowledgments

This work was supported by JSPS Research Fellowships andthe International Advanced Research and Education Organization of Tohoku University.

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