U.S. patent application number 11/356263 was filed with the patent office on 2006-08-17 for ultrasonic diagnosis apparatus.
This patent application is currently assigned to KABUSHIKI KAISHA TOSHIBA. Invention is credited to Eiichi Shiki.
Application Number | 20060184032 11/356263 |
Document ID | / |
Family ID | 19022626 |
Filed Date | 2006-08-17 |
United States Patent
Application |
20060184032 |
Kind Code |
A1 |
Shiki; Eiichi |
August 17, 2006 |
Ultrasonic diagnosis apparatus
Abstract
An ultrasonic diagnosis apparatus comprises an ultrasound probe,
a transmitter (including a transmitting pulse generator and a
transmitting beamformer), a receiver (including a preamplifier and
a receiving beamformer), a CFM processor (including a
moving-element signal extractor and a velocity corrector), a
tomographic image processor, and a display unit. The apparatus
scans a desired section of a subject by transmitting and receiving
an ultrasound pulse to and from the subject, and displays images
obtained by the scanning. The velocity corrector comprises a
pulsation-characterizing-velocity (velocities of a moving element)
calculator, a representative velocity (reference velocity)
calculator, and a corrector to correct the velocities of the moving
element based on the standard velocity. The corrected velocity data
is visualized on display unit. The ultrasonic diagnosis apparatus
makes it possible to display the pulsatility of blood vessels in an
easier and useful way.
Inventors: |
Shiki; Eiichi; (Otawara-Shi,
JP) |
Correspondence
Address: |
OBLON, SPIVAK, MCCLELLAND, MAIER & NEUSTADT, P.C.
1940 DUKE STREET
ALEXANDRIA
VA
22314
US
|
Assignee: |
KABUSHIKI KAISHA TOSHIBA
Tokyo
JP
|
Family ID: |
19022626 |
Appl. No.: |
11/356263 |
Filed: |
February 17, 2006 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
10170573 |
Jun 14, 2002 |
7044913 |
|
|
11356263 |
Feb 17, 2006 |
|
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Current U.S.
Class: |
600/454 |
Current CPC
Class: |
A61B 8/483 20130101;
G01S 15/8984 20130101; G01S 7/52074 20130101; A61B 8/06 20130101;
G01S 15/8988 20130101 |
Class at
Publication: |
600/454 |
International
Class: |
A61B 8/06 20060101
A61B008/06 |
Foreign Application Data
Date |
Code |
Application Number |
Jun 15, 2001 |
JP |
2001-182535 |
Claims
1. An ultrasonic diagnosis apparatus comprising: scanning means for
scanning a subject to be examined while transmitting and receiving
an ultrasound pulse to and from the subject; means for obtaining in
sequence a plurality of velocities of a moving element within the
subject based on reception signals acquired by the scanning
performed by the scanning means; processing means for computing a
reference velocity based on the plurality of velocities obtained
from during a predetermined period of time and correcting each of
the velocities of the moving element using the reference velocity
so that data of the corrected velocities is obtained; and
displaying means for displaying the data so that the data is
updated a plurality of times during the predetermined period of
time.
2. The ultrasonic diagnosis apparatus according to claim 1, wherein
the reference velocity is at least one selected from 1) a mean of
velocities acquired during a predetermined period of time or an
absolute value of the mean thereof, 2) a mean of absolute values of
velocities acquired during a predetermined period of time, 3) an
RMS (Root Mean Squire Value) value of velocities acquired during a
predetermined period of time, 4) a value or an absolute value
thereof, which is calculated by applying any one of an FIR (Finite
Impulse Response) filter, IIR (Infinite Impulse Response) filer,
and a non-linear filter to velocities acquired during a
predetermined period of time or absolute values of the velocities,
and 5) a vectorial mean of velocities acquired from a predetermined
period of time or an absolute value of the mean.
3. The ultrasonic diagnosis apparatus according to claim 1, wherein
the processing means comprises at least one of 1) means for
dividing the velocities of the moving element by the reference
velocity, and 2) means for converting the velocities of the moving
element to values relative to the reference velocity.
4. An ultrasonic diagnosis apparatus comprising: scanning means for
scanning a subject to be examined while transmitting and receiving
an ultrasound pulse to and from the subject; means for obtaining
velocity information about a moving element within the subject
based on reception signals acquired by the scanning performed by
the scanning means; display image producing means for producing an
image to be displayed indicative of a state of pulsation on the
basis of the velocity information; and means for controlling the
display image producing means so that display of the pulsation is
switched on or off on the basis of the velocity information.
5. The ultrasonic diagnosis apparatus according to claim 1, wherein
the processing means has at least one of means for correcting
aliasing resultant from a sampling theorem for the velocities and
means for moderating temporal changes in the data obtained by the
processing means.
6. The ultrasonic diagnosis apparatus according to claim 1, wherein
the scanning means is configured to scan one section of the subject
at the number of frames larger than a value indicated by an inverse
number of a period of time corresponding to an ejection period of
cardiac pulsation.
7. The ultrasonic diagnosis apparatus according to claim 1, wherein
the display means is configured to display a two-dimensional image
of the data obtained by the processing means.
8. An ultrasonic diagnosis apparatus comprising: scanning means for
three-dimensionally scanning one section of a subject to be
examined a plurality of times corresponding to one heartbeat while
transmitting and receiving an ultrasound pulse to and from the
subject; means for obtaining in sequence velocities of a moving
element within the subject based on reception signals obtained
three-dimensionally by the scanning means; processing means for
computing a reference velocity based on the plurality of velocities
obtained from during a predetermined period of time and correcting
each of the velocities of the moving element using the reference
velocity so that data of the corrected velocities is obtained; and
displaying means for displaying at least a three-dimensional image
based on the data so that the image is updated a plurality of times
during the predetermined period of time.
9. The ultrasonic diagnosis apparatus according to claim 8, wherein
the scanning means is configured to three-dimensionally scan the
subject through an electrical scan by means of a two-dimensional
array type of transducer, the velocities of the moving element are
set to a maximum velocity during a predetermined period of time,
and the reference velocity is information that is unchangeable over
cardiac time phases of the subject.
10. The ultrasonic diagnosis apparatus according to claim 8,
further comprising means for obtaining a tomographic image of the
section, wherein the display means has at least one of means for
displaying on the same monitor the tomographic image and the image
of data obtained by the processing means and means for displaying
on the tomographic image the image of data obtained by the
processing means in a superposition manner, the tomographic image
being a two-dimensional image and a three-dimensional image.
11. The ultrasonic diagnosis apparatus according to either of claim
1 or 8, further comprising at least one of means for displaying in
colors the image of data obtained by the processing means, means
for displaying pieces of information formed by combining the data
obtained by the processing means and power information of
scattering echoes from the moving element within the subject, means
for displaying a color bar mapped not only by a hue indicative of a
lower velocity when magnitudes of the data obtained by the
processing means are nearly equal to or less than the reference
velocity but also by other hues indicative of higher velocities as
the magnitude of the data obtained by the processing means becomes
larger than a value nearly equal to the reference velocity, and
means for displaying together both of the data obtained by the
processing means and an aliasing velocity.
12. The ultrasonic diagnosis apparatus according to either of claim
1 or 8, wherein the processing means includes means for obtaining
the velocities of the moving element within the subject pixel by
pixel, the apparatus further comprising means for displaying
together both of a graph showing temporal changes in the velocities
at one or more of the pixels and the image of the data obtained by
the processing means.
13. The ultrasonic diagnosis apparatus according to claim 8,
wherein the processing means comprises processing means for
obtaining factors including a pulsatility index (PI) and a
resistivity index (RI).
14. The ultrasonic diagnosis apparatus according to either of claim
1 or 8, further comprising at least one of means for displaying by
mixture the image of the data obtained by the processing means and
an information of power information of scattering echoes from the
moving element within the subject, means for displaying by mixture
an image of the power information and an image of information
formed by combining the data obtained by the processing means and
the power information, means for displaying together both of a
color bar for the data obtained by the processing means and another
color bar for the power information, means for displaying together
both of a color bar indicative of a combination of the data
obtained by the processing means and the power information and
another color bar for the power information, means for setting an
upper limit and a lower limit on the color bar for the data
obtained by the processing means, and means for displaying at least
one of an upper limit, a lower limit, and an aliasing velocity on
the color bar for the data obtained by the processing means.
15. The ultrasonic diagnosis apparatus according to claim 2,
further comprising means for re-averaging velocities of the moving
element which are equal to or smaller than a value produced by
multiplying the average by "1+.alpha." (.alpha..gtoreq.0) after the
average and for setting to the reference velocity another mean
obtained by the re-averaging.
Description
BACKGROUND OF THE INVENTION
[0001] 1. Field of the Invention
[0002] The present invention relates to an ultrasonic diagnosis
apparatus capable of not only effectively displaying dynamic states
of flow of blood in a subject to be examined, particularly,
pulsation of the flow of blood, but also three-dimensionally
displaying pulsatile flows of blood in the subject.
[0003] 2. Description of Related Art
[0004] An ultrasonic diagnosis apparatus has normally various types
of displaying images, as has been widely known, which can be used
for diagnosis on ultrasound images. Such types of displaying images
include a CFM (Color Flow Mapping) mode used for displaying blood
flow images, as well as tomographic image modes, such as a B-mode,
used for displaying tomographic images.
[0005] Among these modes, the CFM mode is prepared for displaying
two-dimensional blood flow information in real time. In this mode,
generally, flows of blood approaching to the ultrasound probe are
displayed in red on a monitor, while flows of blood going away from
the probe are displayed in blue on the monitor, so that information
about blood flow is visually distinguishable.
[0006] The following describes the principle and an outline of
process for displaying information in relation to blood flow on the
CFM mode. As conventionally well known, an ultrasonic diagnosis
apparatus obtains echo signals sequentially in time by performing
ultrasound scanning at each location (direction) in a subject to be
examined a plurality of times (N-times). Then, from the echo
signals obtained sequentially in time, the apparatus detects
information in relation to velocity and/or scattering power of
blood flow at a desired depth in the scanned region on the basis of
the Doppler technique. That is, scanning the same location in the
subject at predetermined intervals provides Doppler signals
expressed as a quantity of phase shift per unit time of signals
(blood flow signals) reflected from blood cells. The Doppler
signals thus obtained are converted to data of velocity and/or
scattering power of blood flow.
[0007] More precisely, applying quadrature phase detection to the
echo signal at each time of ultrasound scanning with the use of
mixers and LPFs (Low Pass Filters) provides an I (In-Phase) signal
and a Q (Quadrature-Phase) signal, both of which are extracted as
Doppler signals.
[0008] The extracted Doppler signals contain by mixture a reflected
wave signal from objects in motion (moving elements), such as blood
cells, and a second reflected wave signal (called Clatter Signals)
from almost non-moving fixed objects, such as the blood vessel wall
and organ parenchyma. Of these wave signals, the reflected wave
signal from the objects in motion contains a Doppler shift. In
contrast, the reflected signal from the fixed objects hardly
contains a Doppler shift, and is so high in its intensity that the
signal is a dominant in the detected signal.
[0009] Therefore, clatter components representing the reflected
wave from the fixed objects are eliminated through an MTI (Moving
Target Indicator) filter by taking advantage of a difference in the
quantity of Doppler shift, a blood flow Doppler signal can be
efficiently extracted. Then, through analysis of the frequency of
this blood flow Doppler signal, i.e., N-pieces of Doppler data
composed of x.sub.i (I signal) and y.sub.i (Q signal) at each depth
(where i=1, 2, . . . , N), a mean derived from the spectra (i.e., a
Doppler frequency), a dispersion of the spectra, or a reflection
intensity (power) from blood cells can be calculated.
[0010] For this frequency analysis, an autocorrelation function is
normally used. A frequency analysis technique that uses the
autocorrelation function will now be exemplified. As above
described, the blood flow Doppler signal obtained by eliminating
its clatter components at the MTI filter is expressed by a complex
number z.sub.i composed of Doppler data x.sub.i and y.sub.i, each
is N-pieces in number, and expressed by the equation of: z i = x i
+ jy i = a i exp [ j .times. { 2 .times. .pi. .times. .times. f d
.times. T rn .function. ( i - 1 ) + .PHI. ] , ( i = 1 , 2 , ... , N
) ( 1 ) ##EQU1## where a.sub.i is an amplitude, f.sub.d is a
Doppler frequency, T.sub.rn is intervals of transmission of
ultrasound pulses along an arbitrary raster direction, and o is an
initial phase respectively. For the sake of explanatory
convenience, it is assumed that the Doppler frequency f.sub.d is
constant over the N-pieces Doppler data, still maintaining the
generality of the equation.
[0011] According to the above equation (1), the phase rotation of
the complex number z.sub.i per unit time provides a Doppler
frequency f.sub.d. Where a mean complex autocorrelation function
for the N-piece Doppler data is: Z=X+jY=Aexp[j.eta.], (2) the
following equation can be obtained: Z = ( N - 1 ) - 1 .times. i = 1
N - 1 .times. Z i * Z i + 1 = ( N - 1 ) - 1 .times. a i a i + 1
.times. exp .function. [ j .times. { 2 .times. .pi. .times. .times.
f d .times. T rm } ] . ( 3 ) ##EQU2## The Doppler frequency f.sub.d
is therefore expressed as: f.sub.d=(2.pi.)
T.sub.rm.sup.-1tan.sup.-1(Y/X). (4)
X=.SIGMA.(x.sub.ix.sub.i+1+y.sub.iy.sub.i+1)
Y=.SIGMA.(x.sub.ix.sub.i+1-y.sub.iy.sub.i+1)
[0012] By employing this Doppler frequency f.sub.d, the equation
of: V.sub.d=f.sub.dc/(2f.sub.Mcos.theta.) (5) can be obtained, so
that a Doppler velocity V.sub.d is converted using this equation
(5). In this equation (5), c, f.sub.M and .theta. indicate a sound
velocity, the frequency of a reference signal at the mixers, and an
angle formed between an ultrasound beam and each blood flow
(hereinafter referred to as a "Doppler angle"), respectively.
[0013] In the case of a CFM mode, due to difficulty in obtaining
Doppler angles at each position in the space of an image, which
vary position by position therein, the correction of Doppler angles
is omitted from the computation on the foregoing equation (5). In
other words, in the CFM mode, the Doppler velocity V.sub.d can be
calculated based on the equation of: V.sub.d=f.sub.dc/(2f.sub.M)
(6), and is subjected to display in colors. Consequently, where the
Doppler angle is larger, a calculated value becomes smaller than
its original correct velocity, with the result that the Doppler
velocity V.sub.d is subject to display based on color intensities
representing slower velocity states (this is called "angle
dependency").
[0014] Blood flow velocities obtained as described above are
two-dimensionally displayed on a monitor, normally, together with a
B-mode topographic image displayed as a background.
[0015] In recent years, three-dimensional image display in
ultrasonic diagnosis apparatuses has been extensively researched
and developed, and it has been possible to three-dimensionally
display a power image of blood flow. For such a display, for
acquiring three-dimensional data, a hand scanning technique by
which an electronic scanning probe with one-dimensionally arrayed
transducers is used, for example.
To operate this hand scanning, while being electronically scanned
in the direction along the transducer array, an operator moves his
or her hand holding the probe so that the probe is moved to
orthogonal directions to the transducer array direction.
[0016] However, the display on the current CFM mode has encountered
the following problems.
[0017] First, in recent years, as various types of diagnostic
methods have been advanced, there are demands for a display
technique that allows a user to identify a blood vessel as an
artery, portal vein, or vein in a steadier and easier manner. In
particular, to identify a blood vessel as described above by using
an ultrasound wave, it is considered effective to observe pulsation
appearing in blood flow.
[0018] As one conventional examination method for examining
pulsatility of this kind of blood vessel, one display method called
"pulsatility index (PI)" has been known. The PI is an index that
quantifies the extent of change in a blood flow velocity per
heartbeat. Since peripheral circulatory resistance in blood vessels
is reflected in the PI, it is deemed effective for early detection
of dysgenesis of fetuses in the obstetrical department and for
differential diagnosis of tumor in the abdominal part (refer to,
for example, a Japanese Patent Laid-open (KOKAI) Publication No.
H05-317311).
[0019] Other conventional examination methods are provided to
examine the pulsatility of blood vessels, for example. One method
is to display an acceleration of blood flow calculated from two
frames of blood flow velocity data which are adjacent in time and
stored in a frame memory (Japanese Patent Publication No.2768959).
Using this method, information about the pulsation of acceleration
of a blood flow can be added on two-dimensional color flow map
information or three-dimensional display information based on the
CFM mode. Another method is proposed by an ultrasound imaging
method and apparatus that is able to display an image of intensity
of the pulsation appearing in the moving velocity of an echo source
(Japanese Patent Application Laid-open (KOKAI) Publication
No.2000-152935). This method comprises the steps of detecting a
moving velocity of an echo source based on Doppler shifts of
received echoes, detecting intensities of the pulsation appearing
in the moving velocity calculated on moving velocities at the
current and past temporal phases, and producing an image indicative
of the detected intensities of the pulsation. Those methods have
not, however, reached a level of practical use yet.
[0020] Besides, the present CFM mode has a difficulty to clearly
display the pulsatility of blood flow in displaying its power. It
is considered that displaying the velocities of blood flow still
provides distinguishable observation with respect to the
pulsatility. In other words, temporal changes in the colors
indicative of velocities shows that there is pulsatility in a blood
flow to be observed, while no temporal changes in such colors shows
that, there is no pulsatility in the blood flow. However, there are
many cases that make it difficult a clear discrimination of blood
flows even if carefully watched, thereby still lacking practically.
Whichever of the display techniques are chosen, a more convenient
display technique is required to provide the pulsatility of blood
flow.
[0021] Especially in the case of peripheral blood vessels, their
blood flow velocities are relatively lower, amounts of changes in
the velocities showing the pulsatility are also small, even in
arteries. It is therefore considerably difficult to distinguish an
artery from a vein or vise versa on a displayed image. In addition,
displaying the velocity has the problem of the angle dependency if
the Doppler angle is larger, as described before, resulting in that
detected velocities are smaller than their original correct
velocities. It is therefore very difficult to detect the
pulsatility, like the situation in peripheral blood vessels.
[0022] On the other hand, in the foregoing three-dimensional
display, a further advanced display rather than the simple display
of blood vessels is demanded. Such advanced display techniques
include a display technique that has the capability of classifying
the types of blood vessels, such as artery, portal vein, or vein.
For this purpose, it is also considered advantageous that such
display involves pulsatile flows of blood, which requires an
ultrasonic diagnosis apparatus that is able to three-dimensionally
display the pulsatility.
SUMMARY OF THE INVENTION
[0023] The present invention has been made in consideration with
the above problems, and an object of the present invention is to
provide an ultrasonic diagnosis apparatus that is able to
effectively display the pulsatility of blood vessel in a simple and
easy way.
[0024] A further object of the present invention is to provide an
ultrasonic diagnosis apparatus suitable for displaying the
pulsatility of blood vessel in a three-dimensional manner.
[0025] In order to achieve the objects, the ultrasonic diagnosis
apparatus according to the present invention is characteristic of
having, as basic constituents, scanning means for scanning a
subject to be examined while transmitting and receiving an
ultrasound pulse to and from the subject; means for obtaining in
sequence a plurality of velocities of a moving element within the
subject based on reception signals acquired by the scanning
performed by the scanning means; processing means for computing a
reference velocity based on the plurality of velocities obtained
from during a predetermined period of time and correcting each of
the velocities of the moving element using the reference velocity
so that data of the corrected velocities is obtained; and
displaying means for displaying the data so that the data is
updated a plurality of times during the predetermined period of
time.
[0026] In the present invention, it is possible that the processing
means has extracting means for extracting a signal of the moving
element from the reception means.
[0027] In the present invention, by way of example, an
instantaneous velocity can be adopted as a "velocity of a moving
element" (that is, a velocity indicating a characteristic of the
pulsation) within a subject to be examined. In addition, to a
"reference velocity" (that is, a representative velocity), assigned
is at least one selected from 1) a mean of velocities acquired
during a predetermined period of time or an absolute value of the
mean thereof, 2) a mean of absolute values of velocities acquired
during a predetermined period of time, 3) an RMS (Root Mean Squire
Value) value of velocities acquired during a predetermined period
of time, 4) a value or an absolute value thereof, which is
calculated by applying any one of an FIR (Finite Impulse Response)
filter, IIR (Infinite Impulse Response) filer, and a non-linear
filter to velocities acquired during a predetermined period of time
or absolute values of the velocities, and 5) a vectorial mean of
velocities acquired from a predetermined period of time or an
absolute value of the mean. Preferably, the predetermined period of
time is one heartbeat of the subject, a period of time
corresponding to one heartbeat, or a period of time during which
the effects identical to that in one heartbeat are provided.
[0028] In the present invention, the processing means comprises at
least one of 1) means for dividing the velocities of the moving
element by the reference velocity, and 2) means for converting the
velocities of the moving element to values relative to the
reference velocity.
[0029] Preferably, the processing means according to the present
invention has correction means for correcting the aliasing of the
velocities. Further, the present invention may have moderation
means for moderating temporal changes in the data obtained by the
processing means. It is preferred that the scanning means according
to the present invention has scanning means for scanning one
section of the subject at the number of frames larger than a value
indicated by an inverse number of a period of time corresponding to
an ejection period of cardiac pulsation.
[0030] In the present invention, the display means is configured to
display a two-dimensional image based on the data obtained by the
processing means.
[0031] An ultrasonic diagnosis apparatus according to another
aspect of the present invention comprises scanning means for
three-dimensionally scanning one section of a subject to be
examined a plurality of times corresponding to one heartbeat while
transmitting and receiving an ultrasound pulse to and from the
subject; means for obtaining in sequence velocities of a moving
element within the subject based on reception signals obtained
three-dimensionally by the scanning means; processing means for
computing a reference velocity based on the plurality of velocities
obtained from during a predetermined period of time and correcting
each of the velocities of the moving element using the reference
velocity so that data of the corrected velocities is obtained; and
displaying means for displaying at least a three-dimensional image
based on the data so that the image is updated a plurality of times
during the predetermined period of time.
[0032] In the present invention, in the case that the
three-dimensional scan is carried, it is preferred that the
scanning means is configured to three-dimensionally scan the
subject through an electrical scan by means of a two-dimensional
array type of transducer. It is also preferred that the velocities
of the moving element are set to a maximum velocity during a
predetermined period of time. In the case of displaying a
three-dimensional image of information about the pulsatility of the
subject in the present invention, it is preferred that such
information is unchangeable over cardiac time phases of the
subject.
[0033] As another aspect of the present invention, the ultrasonic
diagnosis apparatus further comprises means for obtaining a
tomographic image of the section of the subject, and the display
means is able to display on the same monitor the tomographic image
and the image of data obtained by the processing means. In this
case, it is more effective if the display means is configured to
display on the tomographic image the image of data obtained by the
processing means in a superposition manner. Furthermore, for
three-dimensional display in the present invention, the tomographic
image may be a three-dimensional image. In the present invention,
it is preferable that an image based on the data obtained by the
processing means is depicted in colors.
[0034] As another aspect of the present invention, it is possible
to display pieces of information formed by combining the data
obtained by the processing means and power information of
scattering echoes from the moving element within the subject. In
this configuration, the display can be more effective in cases
where the ultrasonic diagnosis apparatus has means for displaying a
color bar mapped not only by a hue indicative of a lower velocity
when a magnitude of the data obtained by the processing means is
nearly equal to or less than the representative velocity but also
by other hues indicative of higher velocities as the magnitude of
the data obtained by the processing means becomes larger than a
value nearly equal to the representative velocity.
[0035] As another aspect of the present invention, it is preferred
that the image of the data obtained by the processing means and an
information of power information of scattering echoes from the
moving element within the subject are displayed by mixture. It is
also possible to display by mixture an image of the power
information and an image of information formed by combining the
data obtained by the processing means and the power information. In
this case, it is desired that both of a color bar for the data
obtained by the processing means and another color bar for the
power information are displayed together. It may also be possible
to display together both of a color bar indicative of a combination
of the data obtained by the processing means and the power
information and another color bar for the power information. A more
effective example is that setting means is used to set an upper
limit and a lower limit on the color bar for the data obtained by
the processing means, and/or at least one of an upper limit, a
lower limit, and an aliasing velocity on the color bar for the data
obtained by the processing means is displayed.
[0036] As described above, in the ultrasonic diagnosis apparatus
according to the present invention, the scanning means operates to
scan a section to be imaged of a subject through a plurality of
times of transmission and reception of an ultrasound pulse along
the same raster direction of the raster directions required for the
scanning within the subject. The processing means thus operates as
follows. A tissue signal is removed from a reception signal
obtained by the scanning conducted by the scanning means, at every
sample point of the section that has been scanned, so that a blood
flow signal is extracted. Based on this blood flow signal,
velocities of a moving element (i.e., velocities which are
characteristic of the pulsation) and a reference velocity (i.e., a
representative velocity) are figured out. And the velocities of the
moving element are corrected using the reference velocity.
Therefore, under the operations of the display means, data
thus-obtained at individual sample points is depicted as for
example a two-dimensional or three-dimensional image.
[0037] Accordingly, the velocities of a moving element are
corrected based on the reference velocity, which makes it possible
that the pulsatility of slower-speed blood flows, such as flows
passing the peripheral blood vessels, is depicted distinctively.
Further, a Doppler angle dependency is also removed, with the
result that the pulsatility of blood flows is clearly provided even
if a Doppler angle becomes larger. Because the data that has been
subjected to the correction is used for the display, the
pulsatility can be depicted easily and more effectively, compared
to the conventional CFM power images and velocity images. It is
therefore possible to upgrade visibility for the artery, portal
vein and vein, thus contributing to an improved diagnostic
performance.
[0038] In particular, provided that instantaneous velocities are
used as the velocities of a moving element, the display of the
pulsatility becomes excellent, in which dynamic pulsatile changes
are depicted in real time, with the visibility for the pulsatility
enhanced. Suitable for the reference velocity are: a mean of
instantaneous velocities acquired during a predetermined period of
time or an absolute value of the mean thereof; a mean of absolute
values of instantaneous velocities acquired during a predetermined
period of time; an RMS (Root Mean Squire Value) value of
instantaneous velocities acquired during a predetermined period of
time; a value or an absolute value thereof, which is calculated by
applying any one of an FIR (Finite Impulse Response) filter, IIR
(Infinite Impulse Response) filer, and a non-linear filter to
instantaneous velocities acquired during a predetermined period of
time or absolute values of the instantaneous velocities; and a
vectorial mean of instantaneous velocities acquired from a
predetermined period of time or an absolute value of the mean.
[0039] In addition, preferably, the predetermined period of time is
a period of one heartbeat or a period of time corresponding to one
heartbeat. Alternatively, the predetermined period of time may be
any other period of time selected to have the similar effects to
the period of time of one heartbeat. The simplest and most useful
technique for the correction is to use the division. Namely,
utilizing the division allows the pulsatility of slower blood flows
such as peripheral blood flows to be depicted clearly, and the
problem resulting from the Doppler angle dependency can be removed.
Moreover, by correcting the aliasing of velocities, the present
invention is also applicable to faster velocities of moving bodies,
thus making the application more effective.
[0040] In addition, in cases where it is felt that the real-time
display of the pulsatility is hard to observe due to faster
temporal changes thereof on the image, such temporal changes on the
image can be moderated to raise the visibility.
[0041] Further, a specified section of the subject is scanned at
the number of frames larger than a value indicated by an inverse
number of a period of time corresponding to an ejection period of
cardiac pulsation. This scanning makes sure that the ejection
period is traced without failure, whereby the pulsatility can be
depicted in a steadier manner and a diagnostic performance can be
improved. The reason is as follows. The pulsation is caused by the
pumping action of the heart and the ejection period of one cardiac
cycle is characteristic of the pulsation. A blood flow speed at the
vein and the portal vein is almost constant over one cardiac cycle,
while a blood flow speed at the artery increases sharply and then
decreases in the ejection period, and then gradually decreases
until the next ejection period begins. It is therefore very
significant to steadily track the ejection period for detecting the
pulsatility.
[0042] For obtaining a three-dimensional image, the scanning means
is operated to scan a subject in a three-dimensional manner, while
the display means is operated to display a three-dimensional image
based on the data obtained through the scanning. The
three-dimensional scanning carried out by the scanning means
requires a specified section of the subject to be scanned a
plurality of times. This scanning is able to provide both of
velocities of a moving element and a reference velocity at each
section to be scanned, three-dimensional data that has been
corrected, and a three-dimensional image according to the present
invention.
[0043] In cases where such a three-dimensional image based on
information about the pulsatility of a blood flow within a subject
is displayed through the three-dimensional scanning, it is
preferred that a specified section of the subject is scanned
three-dimensionally a plurality of times. This manner will provide
a three-dimensional image in which the pulsatility is depicted with
higher accuracy, whereby such an image is able to contribute to an
enhanced diagnostic performance.
[0044] In order to effectively implement the three-dimensional scan
of the present invention, the most suitable technique is to use a
two-dimensional array type of transducer to three-dimensionally
scan a subject under an electrical staring scanning.
[0045] Furthermore, when a maximum velocity detected during a
period of time is used as velocities of a moving element, a
higher-pulsatility blood vessel, such as the artery, is always
depicted in hues showing higher pulsatility, whilst a
lower-pulsatility blood vessel, such as the portal vein and vein,
is always depicted in hues showing lower pulsatility. Accordingly,
images characteristic of the pulsatility, which are acquired at
each of the cardiac time phases, can be obtained at any time.
Hence, it is still preferable that such an image can be combined
with the foregoing three-dimensional image, wherein two-dimensional
pulsatile images, which become fundamental images for a
three-dimensional image, can be acquired irrelevantly to changes in
the cardiac time phase. Therefore, a three-dimensional pulsatile
image can be constructed easily. In general, when information
indicative of the pulsatility of a blood flow within a subject is
acquired as information irrelevantly to changes in the cardiac time
phase, a three-dimensional pulsatile image can be constructed
easily.
[0046] In the case of displaying the foregoing two-dimensional and
three-dimensional pulsatile images, means for acquiring a
topographic image at a section of a subject can be provided as
well. In such a construction, under the operations of the display
means, both of the tomographic image and an image on the data
obtained by the processing means can be depicted on the same
monitor. This makes it easier to identify the position of a blood
vessel to be observed, thus providing an image of highly improved
visibility. Hence a further enhancement of a diagnostic performance
is possible.
[0047] In particular, it is more effective when the display means
operates to overlay, on the tomographic image, the image on the
data obtained by the processing means. If performing the
three-dimensional display, the tomographic image may be of a
three-dimensional tomographic image, not being limited to a
two-dimensional tomographic image. When the pulsatile image to be
overlaid is depicted in colors, it is sure that the visibility will
be upgraded more, being useful in observing the image. In this
configuration, the data obtained by the processing means may be
combined with scattering power detected from a moving element
within a subject so that an image expressing the combination is
depicted in colors. This display is able to largely raise a visual
effect on the image.
[0048] As another feature, the reference velocity may be a mean of
the velocities that have been detected. In such a construction,
when magnitudes of the data obtained by the processing means are
close to a representative velocity, the above mean is almost equal
to a velocity detected at each of the cardiac time phases of both
the vein and portal vein. In the case of the artery, however, the
above mean is almost equal to a velocity detected at each of the
cardiac time phases other than the ejection period. Considering to
this fact, it is preferable to display a color bar mapped not only
by a hue indicative of a lower velocity when magnitudes of the data
obtained by the processing means are nearly equal to or less than a
reference velocity but also by other hues indicative of higher
velocities as the magnitude of the data obtained by the processing
means becomes larger than a value nearly equal to the reference
velocity. This display of the color bar will make it possible that
the hues assigned to the pulsatility and non-pulsatility is
distinctively distinguished one from the other, resulting in that
the pulsatility can be distinguishably visualized with ease, thus
contributing to a more improved diagnostic performance.
[0049] By the way, breathing, heartbeats, and/or others may cause
an organ to move within a subject. If such a movement happens in
connection with blood vessels, such as peripheral blood vessels or
blood vessels whose Doppler angles are larger, it is conceivable
that blood flow signals are obliged to stop detecting temporarily
or depending on some particular cardiac time phases. In such a
situation, a reference velocity cannot be determined, so that
velocities of a moving element cannot be corrected. In other words,
it is impossible to correct the velocities, but only the detection
of velocities (accordingly, instantaneous blood flow signals) of a
moving object is possible.
[0050] In such a situation, thought the data obtained by the
processing means cannot be subjected the display of an image
thereof (that is, an image indicative of the pulsatility), power
display that represents the existence of a blood flow(s) is at
least possible. Hence, the following display strategy can be given.
Namely, in cases where velocities of a moving element have been
corrected, the data obtained by the processing means is then
subjected to displaying images thereof. By contrast, in cases where
such velocities are no longer corrected, though blood flow signals
have been detected, the corrected velocities are made to undergo
the power display. In other words, an image on the data obtained by
the processing means and a power image are displayed by
mixture.
[0051] This mixing display technique enables the visualization of
all the blood vessels that have been detected. To be specific, the
pulsatility can be depicted toward blood vessels if the velocities
of blood flowing therethrough have been corrected. In that case,
the pulsatility can be visualized at higher vessel detectability,
thus being able to largely improve a diagnostic performance. When
coloring is applied to both of the pulsatile image and the power
image, the visibility to the mixed image can be more raised.
Alternatively, the depiction of a color image concerning
information produced by combining the data obtained by the
processing means and scattering power detected from a moving
element within a subject may be applied to a region at which it has
been possible to correct the velocities. This display technique
will further raise a visual effect on the image.
[0052] In the above display mode, both of a color bar showing the
pulsatility and another color bar showing the power should be put
on the image. The former color bar will help a viewer distinguish
the pulsatility from the other with ease. In addition, in cases
where the depiction of a color image concerning information
produced by combining the data obtained by the processing means and
scattering power detected from a moving element within a subject is
applied to a region at which it has been possible to correct the
velocities, one color bar can be adopted such that hues showing the
pulsatility are taken longitudinally and other hues showing the
power are taken laterally, for instance. This two-dimensional color
bar and the color bar for the power can both be placed on the same
image(s).
[0053] In association with the display of the above color bars,
upper and lower limits of the pulsatility can be put on the color
bars. When such display of the limits is carried out, the
relationship between the hues and degrees of the pulsatility is
clearly specified, thus allowing an observer to easily recognize a
degree of the pulsatility, thus contributing to an enhanced
diagnostic performance. In addition, using means for setting values
to the upper and lower limits on a color bar, an observer can give
proper values to those limits. This will lead to a more effective
display of the pulsatility, whereby an enhanced diagnostic
performance can be expected. Additionally putting an aliasing
velocity on a color bar allows a velocity range to be set
adequately, which results in a steadier detection of the
pulsatility, with the delectability improved.
[0054] As described above, according to one aspect of the present
invention, corrected velocities of blood flow can be displayed two-
or three-dimensionally to provide information about the pulsatility
in a steadier way with the artery, vein, portal vein and others
distinguished one from another. Accordingly the visibility toward
blood flows is greatly raised, both of efficiency and accuracy of
diagnostic examinations are upgraded, thus leading to an improved
diagnostic performance.
[0055] Furthermore, as another aspect of the present invention, a
three-dimensional scan is performed to track the pulsatility of
blood flow through three-dimensional display. In that case, the
three-dimensional scan is carried out with one section scanned a
plurality of times, so that three-dimensional pulsatile data is
provided with higher precision, thereby leading to a greatly
improved diagnostic performance.
[0056] The remaining features of the present invention will be
clearly understood from the following description of preferred
embodiments, which is described together with the accompanying
drawings.
BRIEF DESCRIPTION OF THE DRAWINGS
[0057] In the accompanying drawings:
[0058] FIG. 1 is a block diagram depicting the entire configuration
of an ultrasonic diagnosis apparatus according to a first
embodiment of the present invention;
[0059] FIG. 2 is a schematic block diagram depicting the velocity
corrector;
[0060] FIG. 3 is a schematic block diagram depicting a practical
example of a velocity corrector that divides an instantaneous
velocity by a mean velocity;
[0061] FIGS. 4(a) and 4(b) are time-velocity charts depicting the
operation of an instantaneous velocity extractor, in which FIG.
4(a) is a chart adopting the latest data from among data groups as
an instantaneous velocity and FIG. 4(b) is a chart adopting a
central data from among data groups as an instantaneous
velocity;
[0062] FIGS. 5(a) to 5(c) are views illustrating aliasing with the
use of a vectorial mean value or its absolute value as a mean
velocity data per a period of time of one heartbeat;
[0063] FIG. 6 is a schematic block diagram exemplifying a velocity
corrector which converts velocities characteristic of pulsatility
to representative reference velocities;
[0064] FIG. 7 is a schematic block diagram exemplifying a velocity
corrector that has a function of correcting aliasing
velocities;
[0065] FIG. 8 is a view exemplifying a method of correcting
aliasing velocities;
[0066] FIG. 9(a) depicts a display example of corrected velocities,
and FIG. 9(b) depicts temporal changes in the corrected velocities,
corresponding to FIG. 9(a);
[0067] FIGS. 10(a) to 10(e) illustrate examples displaying color
bars;
[0068] FIGS. 11(a) and 11(b) illustrate an artery displayed on a
conventional velocity mode;
[0069] FIGS. 12(a) and 12(b) are views illustrating the display of
the same artery as shown in FIG. 11, which are displayed using
corrected velocities obtained according to the present;
invention;
[0070] FIG. 13(a) is a chart depicting non-moderated temporal
changes in corrected velocities, FIG. 13(b) is a chart depicting
moderated temporal changes in corrected velocities, FIG. 13(c) is a
block diagram depicting a schematic configuration of a principal
part of a display unit which has a function of moderating changes
in corrected velocities;
[0071] FIG. 14 is a schematic block diagram depicting exemplifying
a display unit that has a velocity corrector;
[0072] FIG. 15 is a schematic block diagram depicting an entire
configuration of an ultrasonic diagnosis apparatus according to a
second embodiment of the present invention;
[0073] FIGS. 16(a) to 16(c) illustrate some examples of a volume
scan;
[0074] FIGS. 17(a) to 17(c) illustrate other examples of the volume
scan;
[0075] FIGS. 18(a) to 18(c) are examples of control sequences for
the volume scan;
[0076] FIGS. 19(a) to 19(c) are illustrations of an electronic scan
with the use of a two-dimensional array probe, which is written in
comparison with other scan methods;
[0077] FIG. 20 is a schematic block diagram practically
exemplifying a velocity corrector that adopts a maximum velocity as
a velocity characterizing the pulsation;
[0078] FIG. 21 is a time-velocity chart depicting the operation of
velocity corrector shown in FIG. 20;
[0079] FIGS. 22(a) and 22(b) are views illustrating display
examples of three-dimensional images;
[0080] FIG. 23 is a time-Doppler velocity chart depicting an
example in which an image indicative of the pulsatility is
unclearly displayed despite that a subject to be examined is
pulsated;
[0081] FIG. 24 is an illustration in which both of a chart
indicating temporal changes in velocities at a pixel pointed by a
marker and a pulsatile image are simultaneously displayed;
[0082] FIGS. 25(a) to 25(c) are various charts indicating temporal
changes in velocities at a pixel pointed by a marker;
[0083] FIG. 26 is an illustration outlining how to calculate a
velocity, which is adopted in cases where a normalized velocity is
used as a corrected velocity;
[0084] FIG. 27 illustrates a mean value of velocity obtained when a
double averaging technique is applied to an artery;
[0085] FIG. 28 illustrates a mean value of velocity obtained when a
double averaging technique is applied to a vein;
[0086] FIG. 29 explains how to use an effective frame rate of
velocity data during a period of time of one heartbeat to determine
whether the calculation of the pulsatility is valid or invalid;
and
[0087] FIGS. 30(a) to 30(d) are views exemplifying display of both
the pulsatility and power in a mixed manner.
DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS
[0088] With reference to the accompanying drawings, embodiments of
an ultrasonic diagnosis apparatus in which the present invention is
implemented will now be described in detail.
First Embodiment
[0089] An ultrasonic diagnosis apparatus according to a first
embodiment of the present invention will now be described with
reference to FIGS. 1 to 14.
[0090] FIG. 1 is a functional block diagram depicting the
configuration of the ultrasonic diagnosis apparatus according to
this embodiment. As shown in FIG. 1, the ultrasonic diagnosis
apparatus according to this embodiment comprises, in addition to an
ultrasound probe 1 (hereinafter simply called "probe") made to
touch the surface of a subject to be examined, a transmitter 2 and
a receiver 3 both electrically connected to the probe 1, a CFM
processor 4 and a tomographic image (B-mode) processor 5 both
electrically connected to the receiver 3, and a display unit 6
electrically connected to both of the processors 4 and 5.
[0091] The probe 1 includes a function of two-way conversion
between an ultrasound signal and an electric signal. One example of
the probe 1 is configured such that an array type of piezoelectric
transducer is linearly set in an array at the distal part of the
probe. This array type of piezoelectric transducer is composed of a
plurality of piezoelectric elements arrayed in parallel, in which
its arrangement direction is assigned to a scan direction. Each of
the plurality of piezoelectric elements forms each channel for
transmission and reception (i.e., called transmission channel and
reception channel).
[0092] The transmitter 2, as shown in FIG. 1, comprises a
transmitting pulse generator 21 for generating a transmission pulse
and a transmitting beamformer 22 for delay-controlling the
transmission pulse from the transmitting pulse generator 21 and
converting the delay-controlled transmission pulse into a drive
pulse applied to each transmission-channel element of the probe 1
for its excitement. Accordingly, the probe 1 is driven, where each
piezoelectric element is able to transmit and receive an ultrasound
pulse to and from the subject.
[0093] The receiver 3, as shown in FIG. 1, comprises a preamplifier
31 placed to each reception channel and a receiving beamformer 32
for performing beam-formation and quadrature phase detection of a
signal received by each preamplifier 31. This configuration makes
it possible that received resigns are delayed and mutually added to
form an echo signal beam-formed in the same direction as that of
the transmission. The echo signal is then subjected to generation
of an I (In-phase) signal and a Q (Quadrature-phase) signal.
[0094] The I and Q signals thus generated make a directional
separation of a Doppler signal possible. That is, those signals can
be used to determine whether a moving element such as blood flow
approaching to the probe 1 or going away from the probe 1. The I
and Q signals (hereinafter simply referred as a "Doppler signal")
are transmitted to the CFM processor 4 and the tomographic image
processor 5, respectively.
[0095] The tomographic image processor 5 produces tomographic image
data of the subject as a B-mode tomographic image from the received
signals obtained through the transmission and reception of
ultrasound pulses. The produced tomographic image data are sent to
the display unit 6.
[0096] The CFM processor 4, as shown in FIG. 1, functionally
comprises, from its signal input side, a moving-element signal
extractor 41 and a velocity corrector 42. As shown in FIG. 1, the
velocity corrector 42 basically comprises calculators 43 and 44 and
a corrector 45. One calculator 43 calculates, as a velocity of a
moving element within the subject, a velocity indicating
characteristics of pulsation (hereafter, the calculator 43 is
referred to as "pulsation-characterizing velocity calculator"). The
other calculator 44 calculates a representative velocity serving as
a reference velocity (hereafter, the calculator 44 is referred to
as "representative velocity calculator") 44. The corrector 45
receives calculation results from both of the calculators 43 and
44, where the corrector 45 corrects the calculated velocity
indicating characteristics of the pulsation on the basis of the
calculated representative velocity.
[0097] In this CFM processor 4, the moving-element signal extractor
41 acquires a blood-flow Doppler signal by removing a clatter
component signal from the echo signal, and then the blood-flow
Doppler signal is sent to the velocity corrector 42. In the
velocity corrector 42, the pulsation-characterizing velocity
calculators 43 calculates, as a velocity of a moving element based
on the present invention, a velocity indicating characteristics of
the pulsation, while the representative velocity calculator 44
calculates a representative velocity serving as a reference
velocity. In addition, the corrector 45 corrects the velocity
indicating characteristics of the pulsation with the use of the
representative velocity, so that data of a corrected velocity is
produced, and sent to the display unit 6.
[0098] The display unit 6 is configured to produce image data in
which, for example, image data of CFM blood-flow corrected
velocities from the CFM processor 4 are overlaid on tomographic
image data from the tomographic image processor 5 and data of a
color bar or others showing the amplitude of the corrected
velocities are placed thereon. The produced image data is displayed
on a monitor. Thus, the corrected velocity data is mapped for
display on the monitor, so that the pulsatility can be visualized
in an easy and simple way without failure. Blood vessels, such as
an artery, portal vein, or vein, can be distinctively displayed,
with the result that visibility for various types of blood vessel
becomes higher, thus improving a diagnostic performance.
[0099] The velocity corrector 42 according to the present invention
will now be exemplified in detail.
[0100] FIG. 2 is a functional block diagram exemplifying a detailed
configuration of the velocity corrector 42 (numbered as a velocity
corrector 42A in FIG. 2). The example of the velocity corrector 42A
shown in FIG. 2 functionally comprises, in addition to the
pulsation-characterizing velocity calculator 43, the representative
velocity calculator 44, and the corrector 45, a velocity calculator
46, a buffer memory 47, and a one-heartbeat-period setting unit
48.
[0101] In this configuration, a blood-flow Doppler signal at each
pixel, which is outputted from the moving-element signal extractor
41 (hereafter this expression will be omitted since the following
is concerning the calculation per pixel) is sent to the velocity
corrector 42A. In the velocity corrector 42A, velocity data are
calculated from the blood-flow Doppler signal at the velocity
calculator 46, and the velocity data that has been calculated are
temporarily stored in the buffer memory 47.
[0102] Velocity data calculated during a period of time of one
heartbeat or its equivalent time (hereafter called "one heartbeat
time") is stored in the buffer memory 47. Whenever new velocity
data comes from the velocity calculator 46, the oldest data in the
buffer memory 47 is removed. Hence, in the buffer memory 47, the
velocity data is updated in real time, while velocity data acquired
during a period of one heartbeat time is held at any time.
[0103] Concurrently with the above data update, the
pulsation-characterizing velocity calculator 43 operates to obtain
a velocity V.sub.cha indicating a specified characteristic of the
pulsation by extracting it from velocity data acquired during one
heartbeat period stored in the buffer memory 47 or from a related
group of velocities. The resultant velocity V.sub.cha is sent to
the corrector 45.
[0104] Meanwhile, the representative velocity calculator 44
calculates a representative velocity V.sub.rep from velocity data
acquired during one heartbeat time in the buffer memory 47, and
send the velocity V.sub.rep to the corrector 45.
[0105] The corrector 45 engages in correcting the velocity
V.sub.cha indicating the specified characteristic of the pulsation
by using the representative velocity V.sub.rep, providing a
corrected velocity V.sub.cmp. The corrected velocity V.sub.cmp is
expressed by: V.sub.cmp=G(V.sub.cha, V.sub.rep), where G is a
function for the correction. The thus-corrected velocity V.sub.cmp
is outputted to the display unit 6.
[0106] Since the velocity corrector 42 provides the representative
velocity V.sub.rep every one-heartbeat time, the representative
velocity V.sub.rep is hardly disturbed by heartbeats. Hence the
corrected velocity V.sub.cmp in which the nature of the
pulsation-characterizing velocity V.sub.cha is accurately reflected
can be obtained.
[0107] In the present embodiment, in response to each time of
update on newly inputted velocity data in the buffer memory 47, the
velocity V.sub.cha indicating a characteristic of the pulsation and
the representative velocity V.sub.rep are also recalculated for the
next update every time new velocity data is received. Consequently,
the corrected velocity V.sub.cmp is also subjected to recalculation
for the next update responsively to the reception of new velocity
data. It is therefore possible for the velocity corrector 42 to
output the corrected velocity V.sub.cmp in real time. The corrected
velocity data is thus excellent in the real-time performance.
[0108] One heartbeat time in the present embodiment is set to both
of the buffer memory 47 and the representative velocity calculator
44 by the one-heartbeat-time setting unit 48. Specifically, the
number L of samples obtained during one heartbeat time is set. The
number L is obtained by dividing a period of time T.sub.HR
equivalent to one heartbeat by a sampling time T.sub.FR for
velocity data. The sampling time shows intervals of time at each of
which velocity data is updated, that is, corresponds to a
reciprocal number of the number of frames.
Therefore, the calculating equation of the number L of samples is
as follows: L=T.sub.HR/T.sub.FR.
[0109] Concerning this equation, a period of time T.sub.HR
corresponding to one heartbeat time is obtained by measuring a
period of time of one heartbeat based on an electrocardiographic
gating signal obtained from a subject with the use of
electrocardiographic gating circuit not shown in FIG. 2. This value
of the time T.sub.HR is transmitted to the one-heartbeat-time
setting unit 48 through a not-shown CPU in FIG. 2. In the case that
an electrocardiographic gating signal is not employed, a period of
time that is approximately equal to one heartbeat is set to the
one-heartbeat-time setting unit 48 with the help of a CPU. A period
of time T.sub.HR corresponding to one heartbeat is, for example,
about one second.
[0110] While the present embodiment has been described based on one
heartbeat, the present invention is not limited to such a case, but
may use a plurality of heartbeats. In the case of a plurality of
heartbeats, the time T.sub.HR is basically interpreted as being
merely repetitions of that obtained in the case of one heartbeat.
However, both a velocity indicating the characteristics of the
pulsation and a representative velocity are usually unchanged from
those values obtained during one heartbeat, thus requiring a larger
capacity of the buffer memory 47 if a plurality of heartbeats are
employed. Accordingly, it is preferable to use one heartbeat. It is
also possible to use an approximate value, which is obtained
through averaging based on a shorter period of time less than a
period of time of one heartbeat, thus requiring a smaller capacity
of the buffer memory 47. In general, it is sufficient if a period
of time equal in effect to a mean over a period of time of one
heartbeat.
[0111] FIG. 3 is a block diagram depicting a schematic
configuration of a velocity corrector 42B, which is another
embodiment of the velocity corrector 42A shown in FIG. 2. In FIG.
3, the velocity corrector 42B adopts an instantaneous velocity
extractor 43A serving as an embodiment of the foregoing
pulsation-characterizing velocity calculator 43, a mean velocity
calculator 44A serving as an embodiment of the forgoing
representative velocity calculator 44, and a divider 45A serving as
an embodiment of the foregoing corrector 45, respectively.
[0112] FIGS. 4(a) and 4(b), which illustrate the operation of the
instantaneous velocity extractor 43A in the above configuration,
are charts depicting a change in velocity V per time t at a certain
one pixel constituting an image (hereafter, all charts to FIGS.
4(a) and 4(b) are shown at one pixel).
[0113] In FIG. 4(a), T.sub.FR indicates a sampling time for
velocity data (T.sub.FR=1/the number of frames), T.sub.HR indicates
a period of time equivalent to one heartbeat, V.sub.1, V.sub.2, . .
. , V.sub.L, V.sub.L+1, V.sub.L+2, . . . , V.sub.L+N indicate
velocity data sampled per time T.sub.FR at a certain pixel, a data
group (0) indicates a group of velocities (V.sub.1, V.sub.2, . . .
, V.sub.L) stored in the buffer memory 47 during a period of time
equivalent to one heartbeat T.sub.HR at a certain time instant, a
data group (1) indicates a group of velocities (V.sub.2, . . . ,
V.sub.L, V.sub.L+1) stored later by one frame than a time phase of
the data group (0) (data groups (2), . . . , (n) in FIG. 4(a) can
be explained in the same way).
[0114] More specifically, in the data group (1), the oldest data in
the data group (0), i.e., V.sub.1 is cleared and the latest data,
i.e., V.sub.L+1 is added. The data in the buffer memory 47 are
updated in this way in turn.
[0115] In FIG. 4(a), the latest data in a data group is adopted as
an instantaneous velocity. For example, a velocity V.sub.L in the
data group (0) and a velocity V.sub.L+N in the data group (n) are
adopted as an instantaneous velocity, respectively. Therefore, in
the instantaneous velocity extractor 43A shown in FIG. 3, these
velocities V.sub.L, V.sub.L+N, . . . are read out in sequence. In
this case, it may be possible to obtain these velocities V.sub.L,
V.sub.L+N, . . . directly from the velocity calculator 46 without
passing through the buffer memory 47.
[0116] Meanwhile, in the mean velocity calculator 44A, a mean value
<V> of velocity data over a period of time of one heartbeat
is calculated. The mean value <V> can be determined as being
any of an mean value of velocities V.sub.1, V.sub.2, . . . ,
V.sub.L obtained at each time phase, the absolute value of a mean
value of velocities V.sub.1, V.sub.2, . . . , V.sub.L obtained at
each time phase, the mean value of an absolute value of velocities
V.sub.1, V.sub.2, . . . , V.sub.L obtained at each time phase, and
an RSM value (Real Mean Square Value) of velocities V.sub.1,
V.sub.2, . . . , V.sub.L obtained at each time phase. Therefore, an
equation for calculating a mean value <V> can be determined
from the following equations: <V>=(V.sub.1+V.sub.2+ . . .
+V.sub.L)/L <V>=|V.sub.1+V.sub.2+ . . . +V.sub.L|/L
<V>=(|V.sub.1|+|V.sub.2|+ . . . +|V.sub.L|)/L
<V>=SQRT{(V.sub.1.sup.2+V.sub.2.sup.2+ . . .
+V.sub.L.sup.2)/L}. In the above equations, L is the number of
samples of velocity data over a period of time of one heartbeat at
each pixel.
[0117] Alternatively, the mean value <V> may be decided as
being a value or its absolute value, which is calculated by
applying an FIR filter, IIR filter, or nonlinear filter to
velocities V.sub.1, V.sub.2, . . . , V.sub.L obtained at each time
phase or their absolute values |V.sub.1|, |V.sub.2|, . . . ,
|V.sub.L|. If defining a function FIL required for the filtering
calculation, equations for the mean value <V> may be
exemplified as follows: <V>=FIL (V.sub.1+V.sub.2+ . . .
+V.sub.L) <V>=|FIL (V.sub.1+V.sub.2+ . . . +V.sub.L)|
<V>=FIL (|V.sub.1|+|V.sub.2|+ . . . +|V.sub.L|)
<V>=|FIL (|V.sub.1|+|V.sub.2|+ . . . +|V.sub.L|).
[0118] Furthermore, a vectorial mean value or its absolute value of
velocities V.sub.1, V.sub.2, . . . , V.sub.L obtained at each time
phase may be set to the mean value <V>. Embodiments of this
setting will now be described with the reference of FIGS. 5(a),
5(b) and 5(c).
[0119] FIGS. 5(a) to 5(c) illustrate aliasing of the Doppler
velocity derived by the sampling theorem. In general, an ultrasound
pulse is transmitted at predetermined intervals T.sub.rn, a blood
flow Doppler signal to be acquired is formed into a discrete signal
consisting of N-pieces of signals that should be mapped at constant
intervals T.sub.rn, if considering one raster. Hence, as shown in
FIG. 5(a), the Doppler velocity is folded back at aliasing
velocities V.sub.a1=.+-.1/(2T.sub.rn) c/2f.sub.M. For example, when
the velocity increases toward its plus (+) direction, the velocity
folds back to the side of -V.sub.a1 at a position immediately after
+V.sub.a1, and then approaches to zero (0) in the reverse direction
along the minus (-) region.
[0120] FIG. 5(b) illustrates the same phenomenon with the use of a
mean complex autocorrelation function Z on N-pieces of Doppler data
defined by the equation (2) described in the section of the
background of the invention. The complex autocorrelation function Z
may be considered as a vector. In other words, in FIG. 5(b), an
increase an a decrease in the velocity means that a vector
corresponding to the complex autocorrelation function Z rotates
around the origin on the complex plane.
[0121] Accordingly, in the case of FIG. 5(b), an angle of rotation
of the complex autocorrelation function Z from +X-axis (strictly
speaking, a phase angle 2.pi.f.sub.dT.sub.rn that is proportional
to the velocity V) corresponds to the velocity V. As a result, the
velocity V=0 is located-on the +X-axis, while the above aliasing
velocity V=.+-.V.sub.a1 are located on the -X-axis. When the
velocity V increases in its plus (+) direction so as to pass the
-X-axis, the velocity turns from its plus (+) value to its minus
(-) value at a boundary of values V=.+-.V.sub.a1.
[0122] Therefore, as shown in FIG.5(c), a mean of velocities
V.sub.1, V.sub.2, . . . V.sub.L obtained at each time phase can be
obtained by adding or averaging complex autocorrelation functions
Z.sub.1, Z.sub.2, . . . , Z.sub.L for each time phase on the same
complex plane. This averaging technique can be understood as a
vectorial averaging technique. This averaging is advantageous in
that it will not be influenced by aliasing, unless the velocity
changes extremely.
[0123] Another method of obtaining a mean of autocorrelation
functions is that the amplitudes of autocorrelation functions Z are
subject to normalization before averaging them. A mean <Z>may
be calculated with the use of one of the equations:
<Z>=(Z.sub.1+Z.sub.2+ . . . +Z.sub.L)/L
<Z>=(Z.sub.1/|Z.sub.1|+Z.sub.2/|Z.sub.2|+ . . .
+Z.sub.L/|Z.sub.L|)/L. (For example, in the case of FIG. 5(c), a
mean <V> of velocities V.sub.1 and V.sub.2 can be-calculated
on <Z>=(Z.sub.1+Z.sub.2)/2.)
[0124] Since the thus-obtained mean <Z> of the
autocorrelation functions is expressed as being
<Z>=<X>+j<Y>, a mean <V> of the velocities
V may be calculated by the following equations, if the previously
defined equations (2) and (4) in the section of the background of
the invention are employed:
<V>=c/(2f.sub.M)(2.pi.T.sub.rn).sup.-1 tan .sup.-1
(<Y>/<X>) <V>=|c/(2f.sub.M)(2.pi.T.sub.rn).sup.-1
tan .sup.-1 (<Y>/<X>)|.
[0125] Concerning the embodiment shown in FIG.3, the above
calculation is implemented by storing autocorrelation functions
Z.sub.1, Z.sub.2, . . . , which are outputted from the velocity
calculator 46, into the buffer memory 47 and by sequentially
reading out the autocorrelation functions Z.sub.1, Z.sub.2, . . .
stored in the buffer memory 47 into the mean velocity calculator
44A. This allows the mean velocity calculator 44A to calculate mean
values <V.sub.0>, <V.sub.1>, . . . of velocity data
included in each of the data groups (0), (1), . . . in turn. Then
the divider 45A is allowed to calculate corrected velocity
V.sub.cmp0, V.sub.cmp1, . . . in turn on the basis of the equations
expressed as follows: V.sub.cmp0=V.sub.L/<V>.sub.0
V.sub.cmp1=V.sub.L+1/<V>.sub.1 . . . .
[0126] In these equations, in lieu of the mean values
<V.sub.0>, <V.sub.1> of the velocity data which are
employed as the denominators, adopting an absolute values of each
of the mean values, the corrected velocities V.sub.cmp0,
V.sub.cmp1, . . . are calculated based on the equations expressed
as follows: V.sub.cmp0=V.sub.L/|<V>.sub.0|
V.sub.cmp1=V.sub.L+1/|<V>.sub.1| . . . . The foregoing
equations for the corrected velocities provide corrected velocity
data with orientation, which has undergone the directional
separation by means of the plus or minus sign of an instantaneous
velocity appearing as a numerator.
[0127] Similarly, in lieu of the corrected velocities V.sub.cmp0,
V.sub.cmp1, the absolute value of each of the velocities
V.sub.cmp0, V.sub.cmp1, . . . can be adopted based on the equations
expressed as follows: V.sub.cmp0=|V.sub.L/<V>.sub.0|
V.sub.cmp1=|V.sub.L+1/<V>.sub.1| . . . , thus providing the
corrected data which indicate their magnitudes only.
[0128] In addition, an instantaneous velocity to be sampled is not
limited to the latest velocity as described before, but any of data
stored in the buffer memory 47 may be usable.
[0129] FIG. 4(b) illustrates an example in which a temporally
median data of each of the stored data groups (0), (1), . . . , (n)
is selected as an instantaneous velocity to be sampled. For
example, in the case of the data group (0), the temporally central
data is data V.sub.L at a median time T.sub.HR/2 of a period of
time T.sub.HR equivalent to one heartbeat. In this case, a time lag
between a mean velocity obtained by the mean velocity calculator
44A and an instantaneous velocity obtained by the instantaneous
velocity extractor 43A is minimized, so that reliability of
corrected velocities is advantageously improved.
[0130] In the foregoing case of FIG. 4(a), the latest data of each
of the stored data groups is selected as an instantaneous velocity
to be sampled. Hence a delay time caused by the processes from data
acquisition through the probe 1 to its display can be minimized,
which provides images with a higher real time performance.
[0131] Although the velocity corrector 42B shown in FIG. 3 has used
the divider 45A functioning as the corrector 45, the present
invention is not limited to this configuration. An alternative
configuration can be provided as shown in FIG. 6, where a velocity
corrector 42C has a mean velocity/reference velocity converter 45B
placed for the divider 45A in FIG. 3. The constituents other than
the mean velocity/reference velocity converter 45B in FIG. 6 (that
is, the velocity calculator 46, buffer memory 47,
one-heartbeat-time setting unit, instantaneous velocity extractor
43A, and mean velocity calculator 44A) are the same in operations
as those shown in the foregoing velocity corrector 42B. The
correction principle on the mean velocity/reference velocity
converter 45B is to divide an instantaneous velocity V by an amount
.zeta. (=.zeta..times.<V>) produced by multiplying a mean
velocity <V> by a certain constant .zeta.. This calculation
lead to a change in a velocity display range for each pixel.
[0132] Specifically, in the case of a calculation technique
according to the embodiment in FIG.6, a value .zeta.<V>
derived by multiplying a mean velocity <V> by a certain
constant .zeta. corresponds to a display range (.zeta. is a
constant independent of the pixels). In addition, a mean velocity
<V> which differ from each other pixel by pixel is multiplied
by a constant .zeta. to produce a velocity range .zeta.<V>
which differs from pixel to pixel. Then an instantaneous velocity V
at each pixel is divided by the velocity range .zeta.<V> at
each pixel. However, in general, the mean velocity/reference
velocity converter 45B operates to covert the instantaneous
velocity V into an amount specific to the velocity range
.zeta.<V> for velocity correction. This correcting example
yields a difference in velocity, which is .zeta. (a constant) times
larger than a velocity corrected by the foregoing dividing
technique, but such a difference can be eliminated when being
allocated to gradations of a color bar later described.
[0133] Since the corrected velocity is originated from a detected
velocity, the calculation of a velocity for characterizing
characteristics of the pulsation and the calculation of a
representative velocity are influenced by aliasing. This aliasing
arises theoretically based on the sampling principle. It is
therefore impossible to erase its effect completely, but the
aliasing can be corrected. One correction technique is the
foregoing vectorial calculation, while there are some other
correction techniques, which will be described below.
[0134] FIG. 7 is a functional block diagram depicting the
configuration of a velocity corrector 42D in which an aliasing
compensation function is additionally implemented. In the velocity
corrector 42D shown in FIG. 7, there is provided the similar
configuration of the velocity corrector 42A shown in FIG. 2, in
which an aliasing corrector 49 is inserted between the velocity
calculator 46 and the buffer memory 47. The other components of the
configuration (that is, the pulsation-characterizing velocity
calculator 43, representative velocity calculator 44, and corrector
45) are the same in their configurations and operations as those in
the embodiment shown in FIG. 2. The aliasing corrector 49 added to
this example is responsible for processing for aliasing correction
that makes use of physical continuity of a velocity change.
[0135] FIG. 8 illustrates an example of processing carried out by
the aliasing corrector 49, which will be explained with the use of
a similar complex autocorrelation function to that shown in FIGS.
5(b) and 5(c). In this example, an assumption is made in such a
manner that, if an interval of time for calculating velocities
V.sub.1, V.sub.2, . . . , V.sub.L at phase times is relatively
shorter than changes in the velocity, velocities at mutually
adjacent time phases are not so far from each other on the complex
plane, whereby aliasing can be corrected. The example shown in FIG.
8 shows a situation where, of two velocities V1 and V2 measured at
mutually adjacent time phases, one velocity V1 is free from
aliasing, but the other velocity V2 is affected by aliasing to give
rise to an erroneous detection of a velocity of -V.sub.2'. In such
a case, an absolute value of a difference between the velocities
V.sub.1 and V.sub.2 is computed, thus establishing an inequality
|-V.sub.2'-V.sub.1|>Va. In this case, calculating an equation of
-V.sub.2'+2Va=V.sub.2 produces |V.sub.2-V.sub.1|<Va. When taking
the physical continuity of velocity into consideration, it is found
that a value of V.sub.2 is more probable than a value of -V.sub.2',
the value of V.sub.2 being adopted.
[0136] Based on the above principle, the aliasing is corrected
by-the aliasing corrector 49. This correction of the aliasing
enables a more broadened range of detectable velocities, improving
accuracy of velocity to be measured. Accordingly, the detectability
for the pulsatility can be enhanced, and a diagnostic performance
is also more improved.
[0137] The corrected velocity calculated as above is displayed on a
screen of the display unit 6. One display example is illustrated in
FIG. 9(a), in which corrected velocities are overlaid on a
tomographic image, but their directions are not separated yet. A
color bar in FIG. 9(a) shows the amplitudes of the corrected
velocities, wherein an exemplified color allocation is such that
corrected velocities having smaller amplitudes are depicted in red
or similar hue thereto, while corrected velocities having larger
amplitudes are depicted in yellow or similar hue thereto.
[0138] FIG. 9(b) illustrates temporal changes in corrected
velocities. This example corresponds to corrected velocities given
by the previously described velocity corrector 42B shown in FIG. 3.
An amplitude "1" of the corrected velocity in FIG. 9(b) shows a
mean velocity, as easily understood from the definitional equation
(for example, V.sub.cmp=V/<V> or others).
[0139] As shown in FIG. 9(b), in the case of a vein or portal vein,
the corrected velocity changes in the neighborhood of the mean
velocity "1." In contrast, an artery shows more drastic changes.
That is, during one cardiac cycle, the corrected velocity increases
sharply over "1" and then decreases sharply in the ejection period,
and then gradually decreases along "1" or thereabouts. In other
words, the ejection period lasts for 200 to 300 ms, while one
cardiac cycle lasts for almost one second. It is therefore
understood that a mean velocity over one cardiac cycle is nearer to
velocities at time phases belonging to a period during which the
velocity gradually decreases.
[0140] Hence, it is preferable that a corrected velocity less in
amplitude than the mean velocity "1" or thereabouts is displayed in
red or reddish hues. In contrast, as a corrected velocity increases
gradually larger in amplitude than the mean velocity "1" or
thereabouts, the display is made to move from yellowish hues to
yellow. Such a display manner is exemplified as shown in FIG. 9(a),
where a blood vessel, such as an artery, with a stronger
pulsatility is depicted in yellowish hues in the ejection period of
one cardiac cycle and in reddish hues in the remaining period. In
contrast, as shown in FIG. 9(a), a blood vessel, such as a portal
vein or a vein, which exhibits a less pulsatility is depicted in
reddish hues throughout one cardiac cycle. This difference in the
depiction of hues makes it possible to distinguishably display the
pulsatility, whereby an operator is able to visually distinguish
differences in the pulsatility in a clear-way.
[0141] In this way, the colors assigned to the color bar is decided
depending on whether or not a corrected velocity is smaller in
amplitude than the mean velocity "1" or thereabouts. When the
corrected velocity is higher in amplitude than this threshold, a
certain hue and its related hues showing higher velocities are
used. When the opposite case to the above comes true, another hue
and its related hues showing lower velocities are used. It is thus
possible to distinguishably assign the hues to both of the
non-pulsatility and pulsatility, thereby providing a visibly easier
distinction to pulsated states, thereby contributing to improvement
in diagnosis.
[0142] FIGS. 10(a) to 10(e) illustrate examples of color bars that
can be displayed.
[0143] FIG. 10(a) illustrates, like the case shown in FIG. 9(b), an
example of a color bar that indicates smaller corrected velocities
in red, for example, and the hue is shifted to yellow, for example,
as the corrected velocity increases. In this embodiment, though the
amplitude of corrected velocities is distinguished by differences
of hues, the present invention is not limited to this way of
display.
[0144] FIG. 10(b) illustrates another example of display of the
color barm, in which the display of power is combined with that of
the corrected velocity shown in FIG. 10(a). In this example, the
larger scattering power of an echo deriving from blood flows, the
brighter a hue to be used in the color bar, and vice versa. This
manner of displaying the scattering power is able to give the
display of vessels a stereoscopic effect, thereby providing a
higher visibility to the vessels to be displayed. The power is
calculated by a not-shown power calculator installed in the
previously described CFM processor 4 in the FIG. 1.
[0145] FIG. 10(c) exemplifies the display of another color bar, in
which directional separation is additionally performed in the
display of corrected velocities shown in FIG. 10(a). In this case,
by way of an example, a flow of blood approaching to the probe 1 is
depicted in warm hues, while a flow of blood going away from the
probe 1 is depicted in cold hues. This directional separation may
make it possible that types of vessels can be distinguished one
from the other in an easier manner.
[0146] FIG. 10(d) also exemplifies another color bar, which is
composed by combining the display of directionally separated
corrected velocities shown in FIG. 10(c) with the display of the
scattering power of echo signals. This display configuration is
able to have the advantages obtained by both of the examples shown
in FIG. 10(b) and 10(c).
[0147] FIG. 10(e) also exemplifies another color bar. In this color
bar, a marker indicative of a threshold is added at the boundary
between the bones for yellow or yellowish hues and red and reddish
hues, which is a display technique adopted by FIG. 10(a) to 10(d)
for displaying the corrected velocities.
[0148] By the way, the present embodiment has adopted the
correction of aliasing of velocities, but it is still impossible to
correct the aliasing perfectly. It is therefore better for an
operator to cope with such an imperfect corrected state of the
aliasing on the monitor of the display unit 6. To realize this
assist, aliasing velocities providing a region of velocities
defined by the sampling principle are depicted on the monitor.
[0149] For example, as previously described FIG. 5(a) to 5(c), the
aliasing velocities are +V.sub.a1 and -V.sub.a1, so those values
are displayed on the monitor. Alternatively, the velocity range may
be broadened by an amount of .DELTA.V in either of (+) or (-)
direction. For example, in cases where the range is broadened in
the (+) direction, the range is changed from a new range of
"-V.sub.a1+.DELTA.V" to "+V.sub.a1+.DELTA.V." Those new values are
preferably depicted on the monitor.
[0150] Since it is usual that an operator knows an approximate
detectable velocity range, the operator may additionally adjust the
aliasing velocity by operating a not-shown button on an operation
panel. The aliasing can therefore be avoided with higher precision.
Hence, images that are more suitable to higher accurate diagnosis
can be obtained.
[0151] Further, since the pulsatility is originated from the
pumping action of the heart, its characteristic aspect is typically
seen during an ejection period of one cardiac cycle. That is, as
described before, while the velocity of blood flow through veins
and portal veins is almost even throughout one cardiac cycle, the
velocity of blood flow through arteries rapidly increases in the
ejection period, and then gradually decreases until the next
ejection period. For detecting the pulsatility, therefore, it is
essential to track the ejection period with certainty. For this
purpose, it is preferable that a cross section of a subject is
scanned at the number of frames higher than a reciprocal number of
a period of time equal to the ejection period of the heartbeat.
This way of scanning enables the ejection period to be tracked
without failure, thus the pulsatility being depicted more
certainly.
[0152] The effect obtained from the display of corrected velocities
according to the present embodiment will now be described in
comparison with that of the conventional one.
[0153] FIGS. 11(a) and 11(b) illustrate the display of an artery
depicted on the conventional velocity mode. FIG. 11(a) illustrates
an embodiment of such display on a monitor. In FIG. 11(a), each
arrow indicates a blood flow velocity (direction and magnitude) of
each of positions .alpha., .beta. and .gamma. located differently
on the artery depicted on the monitor. The magnitudes of the
velocities are V.sub.a, V.sub.b, and V.sub.a, respectively, and the
angles (Doppler angles) formed between ultrasound rasters and blood
flows are .theta..sub.a, .theta..sub.a, and .theta..sub.b,
respectively. In this embodiment, it is assumed that the
relationships of V.sub.a>V.sub.b and
.theta..sub.a<.theta..sub.b are realized.
[0154] FIG. 11(b) shows each blood flow Doppler velocity at each of
the positions .alpha., .beta. and .gamma., i. e., each of temporal
changes in raster-directional signal components "V.sub.acos
(.theta..sub.a)," "V.sub.bcos (.theta..sub.b)," and "V.sub.ccos
(.theta..sub.c)" of blood flow velocities at the positions .alpha.,
.beta. and .gamma.. Since each chart in FIG. 11(b) is concerning
the same blood vessel (artery) shown in FIG. 11(a), temporal
changes in a Doppler velocity at each position should be nearly the
same, if they are measured actually. However, because the velocity
magnitudes are obtained through proportional calculation, it is
found that the magnitudes differ from each other position by
position. Accordingly, the conventional velocity display will give
rise to the situation where the pulsatility on the same artery is
depicted differently from each other position by position, as
described before, despite that the pulsatility is originally the
same on the same artery.
[0155] Practically, as shown in FIGS. 11(a) and 11(b), at such
positions as a narrow vessel of which blood velocity is slower (for
example, at the position .beta.) or a vessel of which Doppler angle
is larger (for example, at the position .gamma.), the resultant
pulsatility is depicted smaller than it is originally. Therefore,
it is actually difficult that the pulsatility at those points is
distinguished visually from blood vessels that exhibit a weaker
pulsatility. In addition, a color bar is displayed which depicts
velocities, for example, in gradually changing hues from "velocity
zero" to positive and negative aliasing velocities. This results in
that the position .alpha. is visualized using a hue showing a
faster velocity and the positions .beta. and .gamma. are visualized
using other hues showing a slower velocity, making the
discrimination of blood vessels more difficult. On top of it, the
range displayed by the color bar and/or how to display on the color
bar are not always suitable for the display of the pulsatility and
the display itself on the color bar is difficult to understand when
viewed. Under such circumstances, it was not easier to correctly
interpret the pulsatility on the conventional velocity mode.
[0156] On the other hand, FIGS. 12(a) and 12(b) illustrate the
display of corrected velocities according to the present invention,
which track and depict the same artery as that shown on the
conventional (shown in FIGS. 11(a) and 11(b)). FIG. 12(a)
exemplifies an image displayed on a monitor, in which there is a
color bar with no directional separation, which is shown similarly
to the foregoing embodiments shown in FIGS. 9(a) and 10(a). FIG.
12(b) illustrates temporal changes in blood-flow Doppler velocities
"V.sub..alpha.cos (.theta..sub..alpha.)," "V.sub.bcos
(.theta..sub.b)," and "V.sub.ccos (.theta..sub.c)" at the
respective points .alpha., .beta., and .gamma., which have been
corrected by means of the previously described velocity correction
technique.
[0157] As shown in FIGS. 12(a) and 12(b), all the corrected
velocities are nearly the same, as long as the same vessel is
concerned. In other words, all of a thinner and
smaller-Doppler-angle portion (like the position .alpha.), a
thinner portion (like the position .beta.), and a
larger-Doppler-angle portion (like the position .gamma.) provides
the same behavior in corrected velocities. That is, the corrected
velocities can be nearly "1" in their gradually-decreasing ranges
and nearly the same value larger than "1" in their strongly
pulsated ranges.
[0158] It can therefore be concluded that the velocity mode of the
present embodiment is independent on flow velocities and Doppler
angles, thus providing approximately equal corrected velocities,
thus depicting almost the same pulsatility, as long as the same
vessel is subjected to display. When such corrected velocities are
subjected to the display with the use of the foregoing color bars,
the display is independent on velocities and Doppler angles, thus
the pulsatility being depicted clearly with largely improved
visibility. In particular, under the velocity display mode
according to the present embodiment, the pulsatility can be
detected even for blood flows, such as blood flows whose velocities
are slower, which are therefore sometimes difficult to be
distinguished from non-pulsated vessels, or another blood flow
subjected to larger Doppler angles. This is able to give images an
effective depiction, thus contributing to greatly improved
diagnosis.
[0159] Besides, as described above, a larger change in the
pulsatility appears during the ejection period lasting 200 to 300
ms at most. Hence, it may feel short for a viewer when visually
observing the pulsatility as it is displayed. In such a case, as
exemplified in FIGS. 13(a) through 13(c), moderating changes in
corrected velocities is effective for making visual observation
easier.
[0160] FIGS. 13(a) and 13(b) exemplify temporal changes in
corrected velocities with no velocity moderation and with velocity
moderation, respectively. In this example, to steadily understand
the pulsatility from the corrected velocities shown in FIG. 13(a),
moderation processing is performed in a way shown in FIG. 13(b).
Specifically, through the moderation processing, the temporal
changes in the corrected velocity are left without any processing
during its rising range, but such changes are moderated so as to
moderately decrease during its descending region, with a sharp fall
being suppressed. This moderation technique allows visibility for
the pulsatility to be enhanced, because there is provided a feeling
of afterimage on images to be displayed.
[0161] Such moderation processing can be performed by, for
instance, the foregoing display unit 6. For realizing the
moderation, the display unit 6 is configured as exemplified in FIG.
13(c). The display unit 6 has, as shown in FIG. 13(c), not merely a
known configuration including a color image memory 61, tomographic
image memory 62, DSC (Digital Scan Converter) 63, and monitor 64
but also a corrected-velocity-change moderator 60 inserted between
the color image memory 61 and the DSC 63.
[0162] In this configuration, the corrected velocity data
transmitted from the foregoing CFM processor 4 is temporarily
stored in the color image memory 61. A plurality of frames of data
is thus stored in the color image memory 61. The
corrected-velocity-change moderator 60 reads out data from the
color image memory 61 to conduct the above-mentioned moderation
processing, and output resultant data to the DSC 63. Apart from
this, the tomographic image data transmitted from the foregoing
tomographic image processor 5 is stored in the tomographic image
memory 62, before being outputted to the DSC 63.
[0163] In the DSC 63, in addition to predetermined image
processing, raster conversion and other types of processing, a
corrected-velocity image that have experienced the moderation
processing in the corrected-velocity-change moderator 60 is
overlaid on a topographic image from the tomographic image memory
62. The resultant image data is then sent to the monitor 64 so that
an synthesized image is depicted thereon.
[0164] In addition, the present embodiment has adopted the CFM
processor 4 that has the function of correcting velocities of
blood-flow Doppler signals. An alternative is that such function is
given to for example the display unit 6, of which configuration is
shown in FIG. 14.
[0165] The configuration shown in FIG. 14 comprises a CFM processor
4B configured in a similar way to the conventional to have the
moving-element signal extractor 41 and the velocity corrector 42
that outputs corrected velocity data. Furthermore, a display unit
6B is also included in the configuration in FIG. 14, and the
display unit 6B is additionally equipped with a velocity corrector
65, in addition to the foregoing known confirmation including the
color image memory 61, tomographic image memory 62, DSC 63, and
monitor 64. The velocity corrector 65, which is located just before
the color image memory 61, is configured similarly to the foregoing
velocity corrector 42A placed in the foregoing CFM processor 4. The
velocity corrector 65 includes a pulsation-characterizing velocity
calculator 66, representative velocity calculator 67, corrector 68,
buffer memory 69, and one-heartbeat-time setting unit 70. This
configuration also enables the display of pulsatility in a similar
way to that described before and provides the advantages identical
to those described before.
Second Embodiment
[0166] Referring to FIGS. 15 to 22, a second embodiment of the
ultrasonic diagnosis apparatus according to the present invention
will now be described. The present embodiment concerns a
configuration in which the correction of the foregoing blood-flow
Doppler signals is applied to three-dimensional display.
[0167] FIG. 15 shows the functional block diagram of the ultrasonic
diagnosis apparatus that will be described below. As shown in FIG.
15, the ultrasonic diagnosis apparatus adopts a two dimensional
array probe 7 as a probe made to touch to the surface of a subject
to be examined.
[0168] As the remaining constituents, the ultrasonic diagnosis
apparatus includes, as shown in FIG. 15, of a transmitter 2A and a
receiver 3A, both of which are electrically connected to the
two-dimensional array probe 7. The apparatus further includes a CFM
processor and a topographic image processor 5, both of which are
electrically connected to the receiver 3A, and a display unit 8
electrically connected with both the processors 4 and 5. Of these
constituents, the transmitter 2A includes, in addition to the
transmitting pulse generator 21 and the transmitting beamformer 22,
which are identical to those in FIG. 1, a scan controller 2C to be
added. The receiver 3A includes the preamplifiers 31 and the
receiving beamformer 32, like the configuration in FIG. 1. The CFM
processor 4 includes, like the configuration in FIG. 1, the
moving-element signal extractor 41 and the velocity corrector 42.
Furthermore, the display unit 8 includes, in addition to the DSC 83
and monitor 84 that have been described in FIG. 13(c), a
three-dimensional color image memory 81 and a three-dimensional
tomographic image memory 82. (The constituents similar or identical
to those described before will be omitted from being explained
below or will be explained just briefly.)
[0169] The two-dimensional array probe 7 has a function of two-way
conversion between ultrasound signals and electric signal. To be
specific, the probe 7 has a two-dimensional array type of
piezoelectric transducer two-dimensionally placed at the tip. The a
two-dimensional array type of piezoelectric transducer is formed by
mapping a plurality of piezoelectric elements two-dimensionally,
that is, like a matrix, so that a ultrasound beam signal can be
scanned three-dimensionally, including the longitudinal, lateral,
and oblique directions. A plurality of piezoelectric elements
constitute transmitting and reception channels, respectively.
[0170] Hence, the two-dimensional array probe 7 is allowed to scan
a plurality of sections located within a subject, as sections to be
scanned are changed (i.e., a "volume scan" is performed), with the
result that three-dimensional data (i.e., volume data) can be
acquired. The volume scan is exemplified as various forms, as shown
in FIGS. 16(a) to 16(c) and FIGS. 17(a) to 17(c).
[0171] FIG. 16(a) explains a scan technique by which a section to
be scanned is shifted along a perpendicular direction to the
section during its scanning operation. Meanwhile, FIG. 17(a)
explains another scan technique used in such a manner that a
section to be scanned is shifted to rotate about its central axis
during its scanning operation. When either technique is used, it is
required that a single section be scanned a plurality of times,
preferably, a plurality of times during a period of one heartbeat.
This scanning is one of the characteristic inherent to present
invention.
[0172] One scanning example is shown in FIGS. 16(b) and 17(b),
respectively, wherein an ultrasound beam scans a certain section
during a period of one heartbeat, and then the ultrasound beam is
shifted by a width equal to a beam thickness to scan the next
section in the same manner. Another scanning example is shown in
FIGS. 16(c) and 17(c), respectively, wherein an ultrasound beam
scans a certain section for one frame of data, and then the beam is
shifted slightly by a width smaller than a beam thickness to san
the next section for one frame of data in the same manner as above.
Either scanning technique permits any location, which is
arbitrarily selected in the scanned three-dimensional space, to be
scanned during a period of one heartbeat. This volume scan makes it
possible ultimately to calculate corrected velocities in each
section, thus making it possible ultimately to display the
pulsatility in the three-dimensional manner.
[0173] The scanning techniques shown in FIGS. 16(b) and 17(b) are
superior in respect of accuracy for calculating corrected
velocities, since each section to be scanned is stationary during a
period of one heartbeat. On the other hand, the techniques shown in
FIGS. 16(c) and 17(c) are advantageous in that corrected velocities
can be calculated specially with fineness, because sections to be
scanned are consecutively located.
[0174] The scan controller 23 of the transmitter 2A (refer to FIG.
15) can realize these scanning technique, in which the controller
23 is operated to control scanning toward each section to be
scanned. Sequences for such scan control are exemplified in FIGS.
18(a) to 18(c).
[0175] Of these, FIG. 18(a) shows an example of a scan sequence
used for a conventional one-dimensional array type of probe. This
sequence explains that a plurality of rasters composing the same
section to be scanned (refer to "raster 1, raster 2, . . . , raster
J" in FIG. 18(a)) are repeatedly scanned for each frame of data
(refer to "1.sup.st frame, .sub.2.sup.nd frame, . . . " in FIG.
18(a)).
[0176] In contrast, FIGS. 18(b) and 18(c) explain examples of scan
sequences directed to the volume scan performed with the aid of the
two-dimensional array type of probe according to the present
embodiment.
[0177] Of these examples, the sequence exemplified in FIG. 18(b) is
directed to the foregoing scanning techniques shown in FIGS. 16(b)
and FIG. 17(b). Practically, as shown in FIG. 18(b), with respect
to a certain one section located within a spatial region to be
subjected to the volume scan (refer to "section 1" in FIG. 18(b)),
a plurality of rasters composing the section (refer to "raster 1-1
(meaning that a raster 1 on the section 1), raster 1-2, . . . ,
raster 1-J" in FIG. 18(b)) are scanned to produce data for one
frame, and then the same scanning to the plurality of rasters of
the same section are repeatedly performed I-times (i.e., for
I-pieces of frames), which leads to acquisition during a period of
one heartbeat. After completing the scanning at the foregoing
section, the ultrasound beam is controlled to scan the next section
(refer to "section 2" in FIG. 18(b)) is shifted stepwise from the
previous one by a width equal to a beam thickness. At the section
2, the scanning is also performed similarly to the above during a
period of one heartbeat. The same scanning manner is repeatedly
applied to each of the remaining sections located in the spatial
region realized by the volume scan.
[0178] In addition, the sequence exemplified in FIG. 18(c) is
directed to the foregoing scanning techniques shown in FIGS. 16(c)
and FIG. 17(c). Practically, as shown in FIG. 18(c), first of all,
a plurality of rasters (refer to "raster 1-1, raster 1-2, . . . ,
raster 1-J" in FIG. 18(c)) constituting a certain one section
located within a volume region to be scanned (refer to "section 1"
in FIG. 18(c)) undergo scanning for one frame of data by an
ultrasound beam. Then, the ultrasound beam is controlled to scan
the next section (refer to "section 2" in FIG. 18(c)) slightly
shifted from the previous one by a width less than a beam thickness
(for instance, a width equal to 1/I of a beam thickness). The next
section (i.e., "section 2") is then subjected to scanning for one
frame of data in the same manner as above. Hereafter, as sections
to be scanned are shifted little by little, the scanning for
"section 3, . . . , section I, section I+1, . . . " is conducted
consecutively in turn. A period of time for one heartbeat is
required to scan I-piece sections. The data acquired from the
sections ranging from the section 1 to the section I is used to
calculate corrected velocities, so that one section on which the
corrected velocities are mapped is produced. After a delay of time
for one frame, the data acquired from the sections ranging from the
section 2 and the section I+1 is used to calculate corrected
velocities. Hence, the next section on which the corrected
velocities are mapped is produced. Hereafter, in the same manner as
above, other sections on each of which corrected velocities are
mapped are consecutively produced with a delay of time for one
frame left therebetween, whereby the plural sections mutually
adjacently allocated in time are produced.
[0179] For performing the volume scan, a parallel simultaneous
reception technique may be used which enables simultaneous
reception along a plurality of directions in response to one time
of transmission for scanning a section. For instance, this
technique can be applied to simultaneous scans toward a plurality
of sections, providing various advantages including a shortened
time for scanning a volume region. As long as each section is
scanned during a period of time for one heartbeat, the parallel
simultaneous reception technique can be implemented into various
types of scanning techniques.
[0180] Because exemplified in the present embodiment is an
electronic scan that uses the two-dimensional array probe 7,
advantages derived therefrom will now be explained compared to the
other scanning methods (that use a one-dimensional array type of
probe).
[0181] FIG. 19(a) illustrates a conventionally known scan
technique, called hand scan, by using an electronic scan probe
(one-dimensional array type of probe) with a one-dimensionally
arrayed transducer. An operator holds the probe by hand, and moves
the probe in a perpendicular direction to the sections to be
scanned every time each section has been scanned in a usual
scanning manner. As stated before, a consecutive operation to
positionally shift sections to be scanned with data acquired during
a period of time for one heartbeat involves a heavier difficulty,
thus being not practical.
[0182] Further, FIG. 19(a) is another illustration explaining how
to scan each section, which has been known as well. That is, a
conventional electronic scan probe (one-dimensional array type of
probe) with a one-dimensionally arrayed transducer is mounted to a
known probe-moving unit with a guide mechanism and rotation
mechanism which use guide rods. The probe-moving unit is operated
to slowly move the probe so that each section is scanned during a
period of time for one heartbeat for acquisition of volume data.
This unit allows each section to scanned over a period of time for
one heartbeat. However, the unit should be large-scale and the
operations between scans and the prove movements should be in
synchronism with each other, thus complicating the control, thus
being not practical as well, as explained about FIG. 19(a).
[0183] In contrast, FIG. 19(c) is an illustration to explain an
electronic scan with the use of the two-dimensional array type of
probe 7 according to the present embodiment. When using this scan
technique, it is sufficient for an operator to only fix the probe 7
on a desired region of a subject to be examined. This operation
makes it possible that volume data is automatically acquired and
depicted with higher accuracy and the pulsatility is
three-dimensionally depicted with precision in a simpler
manner.
[0184] Based on the data acquired by the two-dimensional array type
of probe 7, velocities are corrected. For example, maximum
velocities can be employed as the corrected velocities, whereby it
is easier to perform three-dimensional reconstruction. The velocity
corrector 42 responsible for the foregoing correction based on the
maximum velocities is exemplified in its construction in FIG.
20.
[0185] A velocity corrector 42E shown in FIG. 20 comprises, in
addition to the velocity calculator 46, buffer memory 47, mean
velocity calculator 44A, divider 45A, and one-heartbeat-time
setting unit 48, as similarly to FIG. 3, a maximum velocity
detector 43B serving as the pulsation-characterizing velocity
calculator. FIG. 21 exemplifies the operation at the maximum
velocity detector 43B.
[0186] In the similar way to that in FIGS. 4(a) and 4(b), changes
in velocity V over time t at a certain one pixel are shown in FIG.
21. In FIG. 21, T.sub.FR indicates a sampling time for velocity
data (T.sub.FR=1/the number of frames); T.sub.HR indicates a period
of time for one heartbeat; V.sub.1, . . . , V.sub.I, . . . ,
V.sub.L, V.sub.L+1, . . . , V.sub.1I . . . indicate velocity data
sampled per sampling time T.sub.FR at a certain pixel; a data group
(0) indicates a group of velocity data (V.sub.1, V.sub.2, . . . ,
V.sub.L) stored in the buffer memory 47 during one heartbeat time
T.sub.HR at a certain time instant; and a data group (1) indicates
a group of velocity (V.sub.2, . . . , V.sub.L+1) stored one frame
later than a time phase at which the data group (0) is obtained.
The remaining data groups (2), (3), (4) . . . are also obtained in
the same way.
[0187] In buffer memory 47, the data update is carried out in such
that the oldest data in the data group, that is, velocity data
V.sub.1, is removed, while the latest data, that is, V.sub.L+1, is
added. This updating technique is repeated at intervals in
turn.
[0188] In response to the operations in the buffer memory 47, the
maximum velocity detector 43B operates to read out a maximum
velocity from each of the groups of velocity data. Practically, in
the case of the example shown in FIG. 21, a velocity V.sub.I is
read out of each of the data groups (0) to (2) and a velocity
V.sub.II is read out of each of the data groups (3) and (4). The
read maximum velocities are sent in sequence to the divider 25 45A.
Because each of the groups of velocity data consists of velocities
acquired during one heartbeat time T.sub.HR, every data group
always includes data of a maximum velocity appearing during a
period of tome for one heartbeat. In addition, each of the groups
of velocity data to be stored is updated at any time, whereby the
latest velocity is always new.
[0189] In parallel, as described before, the mean velocity
calculator 44A is engaged in the consecutive calculation of each
mean value <V>.sub.0 (<V>.sub.1, <V>.sub.3,
<V>.sub.4, . . . ) of each of the groups of velocities,
whereby the resultant values are routed to the divider 45A.
[0190] Thus, the divider 45A receives both of a maximum velocity
V.sub.I (V.sub.II, and others) from the maximum velocity detector
43B and a mean velocity <V>.sub.0 (<V>.sub.1,
<V>3, <V>.sub.4, and others) from the mean velocity
calculator 44A, and then calculates consecutively a corrected
velocity V.sub.cmp0 (V.sub.cmp1, V.sub.cmp2, V.sub.cmp3, and
others) using themes values that have been received, on the basis
of the equations expressed as follows:
V.sub.cmp0=V.sub.I/<V>.sub.0
V.sub.cmp1=V.sub.I/<V>.sub.1
V.sub.cmp2=V.sub.II/<V>.sub.2
V.sub.cmp3=V.sub.II/<V>.sub.3 . . . .
[0191] In these equations, in lieu of mean velocities that are the
denominators, the absolute values of the mean velocities may be
adopted, whereby the corrected velocities can be calculated on the
following equations. V.sub.cmp0=V.sub.I/|<V>.sub.0|
V.sub.cmp1=V.sub.I/|<V>.sub.1|
V.sub.cmp2=V.sub.II/|<V>.sub.2|
V.sub.cmp3=V.sub.II/|<V>.sub.3|
[0192] Thus, the data of corrected velocities that have been
oriented can be obtained, where each of which is directionally
separated on the positive and negative signs of instantaneous
velocities that are numerators.
[0193] In addition, corrected velocities may be obtained as their
absolute values by the following equations:
V.sub.cmp0=|V.sub.I/<V>.sub.0|
V.sub.cmp1=|V.sub.I/<V>.sub.1|
V.sub.cmp2=|V.sub.II/<V>.sub.2|
V.sub.cmp3=|V.sub.II/<V>.sub.3| so that the obtained values
indicate only magnitudes of the corrected velocities. This
calculation is identical to that previously shown on FIG. 3.
[0194] This calculation enables the display of only maximum
velocities that have been corrected. For instance, when applied to
the foregoing example shown in FIG. 5, the artery is always
depicted in yellow showing faster velocities, while the portal vein
and vein are always depicted in red showing slower velocities.
Therefore, an image, which is like a still image, can be displayed.
In general, if pieces of information independent of changes in the
cardiac time phase is used as pieces of information in relation to
the pulsatility, an image, which can be seen like a still image, is
displayed. This kind of image is especially effective when it is
desired to observe still images, preserve data of still images,
produce three-dimensional images, and others.
[0195] The maximum velocities that have been corrected are
constantly updated in real time, without being frozen. Thus, even
if a blood vessel is slightly moved in the subject's body due to,
for example, a soft breathing, the blood vessel on an image can
also be moved and depicted, still providing effective information
that tracks the vessel. Accordingly, using such maximum velocity
data that has been corrected makes it possible to provide pulsatile
images that are excellent in visibility, preservation, and
real-time performance, and suitable for constituting
three-dimensional images. As a result, a diagnostic performance can
be improved further.
[0196] The data of corrected velocities calculated by the velocity
corrector 42E is then sent to the display unit 8, as described
before (refer to FIGS. 15 and 20). The display unit 8 has, as
described in FIG. 15, the three-dimensional color image memory 81
and the three-dimensional tomographic image memory 82, both of
which serves as image memories and are able store three-dimensional
data therein. The foregoing data of corrected velocities is thus
stored in the three-dimensional color image memory 81, while
tomographic image data two-dimensionally or three-dimensionally
acquired and processed is stored in the three-dimensional
tomographic image memory 82.
[0197] When the foregoing volume scan to acquire volume data is
completed, the three-dimensional data of corrected velocities
stored in the three-dimensional color image memory 81 and the
two-dimensional or thee-dimensional tomographic image data stored
in the three-dimensional tomographic image memory 82 are both read
out. The read-out data are then subject to display processing
according to a format specified and inputted by an operator through
a not-shown operation panel, with the result that a
three-dimensional image is depicted on the monitor 84 in
cooperation with the DSC 83. Examples of display of such
three-dimensional images are shown in FIGS. 22(a) and 22(b).
[0198] FIG. 22(a) illustrates an image formed by combining a
two-dimensional topographic image with a three-dimensional
pulsating vessel image, whilst FIG. 22(b) illustrates an image
formed by combining a three-dimensional tomographic image with a
three-dimensional pulsating vessel image.
[0199] Accordingly, the above characteristics of image display can
be applied to proper use originated from necessity. A pulsating
vessel can therefore be displayed three-dimensionally, during which
time a user is allowed to identify a desired position or observe a
lesion on a topographic image. This makes it possible that both of
the artery and the vein are observed with improved vessel
continuity, greater visibility, convenience, and clearer
distinction, whereby the efficiency and accuracy of each
examination can be improved to a greater extent.
[0200] As explained above, the foregoing first and second
embodiments permit the pulsation of blood vessels to be tracked,
thus being in principle depicted distinguishably between the artery
and the vein. However, it is clear that there is no one-to-one
correspondence as to the pulsatility between the artery and vein,
and it is conceivable that some exceptions may occur.
[0201] For example, an inferior vena cava indicates to some extent
the pulsatility. In such a case, even a vein may provide an image
in which the pulsatility appears to a small extent. In contrast, a
Doppler image shown in FIG. 23 may be provided, where there are
smaller changes in velocity compared to a mean velocity over one
heartbeat, even though a blood flow with a tumor exhibits the
pulsatility. In such a case, there is a possibility that the image
does not provide a clear pulsatility of a blood vessel to be
observed, despite the real existence of the pulsation thereof.
[0202] Countermeasures for improving the above situation will now
be described on examples preferably shown in FIGS. 24 and 25.
[0203] In the case of such examples, first of all, the calculation
for velocity at individual pixels is carried out by the CFM
processor to obtain velocities at the pixels, like the
conventional. Then, as shown in FIG. 24, a maker is placed on a
pulsatile image to specify a certain pixel. Responsively to this
specification, temporal changes in velocity at the specified pixel
(that is, a velocity-time chart) are depicted together with the
pulsatile image.
[0204] This display technique allows spectrum Doppler data to be
obtained from the same scanned data as those used for display the
pulsatility, thus preventing the number of frames from being
lowered. In contrast, the conventional parallel display of CFM
images and spectrum Doppler data require to be scanned separately
and dedicatedly from each other, resulting in that the number of
frames which can be used for CFM images is decreased down to half
of the frames. The advantage of keeping the number of frames, which
can be obtained by the present embodiment, is thus strictly
significant if it is desired that an ejection period be securely
tracked for displaying the pulsatility. Moreover, another advantage
of the present embodiment is that temporal changes in velocities
detected at a plurality of positions are displayable, although the
conventional parallel display of CFM images and spectrum Doppler
data allows Doppler spectrum data to be displayed at only one
position. Thus, the present embodiment is able to provide a
remarkably improved usage. Of course, in the present embodiment,
Doppler spectrum data at a plurality of positions can be displayed
in real time in synchronism with displaying pulsatile images.
Additionally, only a chart showing velocities versus time or only
pulsatile images can be frozen.
[0205] A modification is that a marker used in the present
embodiment is variable in its size. Thus, spatial averaging of data
enclosed by the size-changeable marker is able to stabilize
velocity data that will be produced. In the present embodiment,
plotting a velocity at each frame forms charts showing the
relationship between velocities versus time, which are exemplified
in FIGS. 25(a) to 25(b). As shown therein, how to plot velocities
can be changed. For instance, a velocity obtained at each frame is
plotted with the use of a dot. Alternatively, such dots can be
connected with polygonal lines or a curved line. Another
modification is that an electrocardiographic gating technique is
used to obtain velocities as the cardiac time phase is shifted
gradually and velocities obtained during several heartbeats are
overlaid one on another, thus providing a chart with closer
intervals of time. Still, another modification is concerned with
the calculation of velocities, in which a mean velocity at each
pixel, which is usually obtained for only the display, is combined
with the dispersion of a velocity at the pixel, in order to obtain
a pseudo maximum velocity for display. Moreover, the foregoing
embodiment is suitable for real-time display, but this is not a
definite list. By way of example, the display of the pulsatility
may be frozen, before a velocity-time chart at a specified pixel is
displayed, if velocity data is once stored in a memory.
[0206] The above various modifications will upgrade images that
depict pulsatile flows, making it possible to detect pulsatility in
a more sure fashion, thus further improving a diagnostic
performance.
[0207] The foregoing pulsatile flow display can be applied to the
following first to third examples.
[0208] 1) The first application example is concerned with combining
the foregoing pulsatile flow display with a known broad-band
transmission technique (refer to, for example, Japanese Patent
Laid-open (KOKAI) publications Nos. 2000-342586 and 2001-269344).
This "broad-band transmission" technique, which is carried out
under the power Doppler mode, is a way of transmitting an
ultrasound pulse consisting of one or two burst waves, instead of
transmitting an ultrasound pulse consisting of four to eight burst
waves. The one or two burst waves are equivalent in the number of
burst waves to that under the B-mode, thus providing a broad-band
ultrasound pulse. This "broad-band transmission" technique is able
to provide range discrimination essentially equivalent to the
B-mode (for example, refer to a Japanese Patent Laid-open (KOKAI)
Publication No.2000-342586).
[0209] One example of the broad-band ultrasound pulse is an
ultrasound pulse of which number of burst waves is less than three.
By using this ultrasound pulse, it becomes possible to obtain blood
flow images of high spatial resolution with no or almost no
blooming, so that diagnostic performance is further improved (refer
to a Japanese Patent Laid-open (KOKAI) Publication No. 2000-342586,
for example). Computing the reciprocal number of a transmission
frequency of a transmission ultrasound pulse produces one cycle of
the transmission pulse. An ultrasound pulse of which duration is
equal to one cycle of a transmission pulse is called a "pulse of
one burst wave" and an ultrasound pulse of which duration is equal
to two cycles of a transmission pulse is called a "pulse of two
burst waves." Thus, an ultrasound pulse of which duration is equal
to M-cycles of a transmission pulse is called a "pulse of M-piece
burst waves." One burst waves, two burst waves, . . . , M-piece
burst waves are called the number of burst waves.
[0210] Accordingly, the present application example will provide a
further improved resolution performance, in addition to the effects
derived from the foregoing pulsatile flow display.
[0211] 2) The second application example concerns an improved way
of calculating the foregoing reference velocity.
[0212] FIG. 26 exemplifies temporal changes in velocity (i.e.,
changes in velocity in the frame direction) at a certain one pixel
on a colored map providing a blood flow image, which is adopted to
explain a formula for calculating the foregoing corrected
velocities. In this example, an example is provided without loosing
the generality such that the reference velocity is given as a mean
of absolute values of velocities over one heartbeat time and
corrected velocities are given as "corrected velocities=absolute
values of instantaneous velocities/a mean of absolute values of
velocities acquired over one heartbeat time." As described above,
the corrected velocities are obtained by dividing velocity data by
a velocity at the same pixel, thereby dissolving the drawback from
the Doppler angle dependency. Thus the pulsatility at a blood
vessel can be detected more clearly. In this case, normally, a mean
is calculated through one time of averaging over all instantaneous
velocities acquired over one heartbeat time.
[0213] In addition to the foregoing one-time averaging technique, a
two-time averaging technique can be provided as the technique for
calculating a mean according to the present application example,
which is shown in FIG. 27 (for the artery) and in FIG. 28 (for the
vein). The two-time averaging technique is used to average all
instantaneous velocities acquired over one heartbeat time so that a
mean M1 is figured out, to average only instantaneous velocities
equal to or less than a value (=M1.times.Th-value) obtained by
multiplying the mean M1 by a predetermined value (corresponding to
Th-value; for instance, Th-value=1+.alpha.) so that a mean M2 is
figured out through the second averaging, and to adopt the mean M2
for the mean M1 as the denominator in the foregoing calculating
formula. In other words, instantaneous velocities larger than the
value of M1.times.Th-value are excluded from the calculation of the
mean M2.
[0214] This two-time averaging technique is advantages as follows.
1) In the artery shown in FIG. 27, compared to the mean M1 obtained
from the one-time averaging technique, the pulsatility is enhanced
correspondingly to a decrease in the mean M2 derived from the
second averaging calculation. In contrast, in the case of the vein
shown in FIG. 28, there is almost no difference between the means
M1 and M2, so that the-vein is depicted as a non-pulsatile flow. 2)
The mean M2 figured out by the second averaging is sensuously more
closer to a mean obtained thorough manual tracing on the image
shown in FIG. 27, thus being easier to understand. 3) Even only an
ejection period has been subjected to the detection, the
pulsatility is easier to display. Therefore, the artery and vein
are more clearly distinguished one from the other; whereby a
diagnostic performance can be improved more.
[0215] 3) The third application example is directed to a display
mode, in which both of the foregoing pulsatile flow display and a
known power display are carried out by mixture.
[0216] In the case that an organ is moved by breathing, heartbeats,
or others, or a blood vessel to be observed is a peripheral blood
vessel or a blood vessel with the large Doppler angle, it happens
that the blood flow signal may not be detected temporarily or at
some cardiac time phases. Resultantly, the reference velocity may
not be determined, which make it impossible to correct velocities
of a moving element.
[0217] To be specific, in cases where an organ moves, any one pixel
on a blood vessel may not be present at the same location during
one heartbeat time. Hence, all the velocity data of a moving
element which should be gathered during one heartbeat time are not
always fully acquired, thereby lacking data.
[0218] In addition, in the case of a peripheral blood vessel, a
velocity of blood is slow, so that a possible Doppler signal
frequency to be detected becomes lower. In the case of a blood
vessel having a larger Doppler angle, a Doppler signal frequency to
be detected is also obliged to be lower due to its angle
dependency. In such cases, an MTI filter operates to eliminate a
signal of which Doppler frequency is lower. This means that, for
example, it is possible to detect Doppler signals during an
ejection period at the artery, while it is difficult to fully
detect Doppler signals during a diastole thereat. If such an
occasion occurs, it is impossible to prepare for all velocity data
of a moving element during one heartbeat time, thus lacking
data.
[0219] The present application example is to cope with such a
situation. For this purpose, this example is configured such that
the reference velocity can be determined even when velocity data of
a moving element lacks to some extent. Hereinafter, this will be
described about one pixel on an image.
[0220] As shown in FIG. 29, it can be assumed that there are 20
frames in one heartbeat time. By way of example, if velocities of a
moving element can be detected in 70% or more of all the 20 frames,
i.e., in 14 frames or more, it can thus be determined that a
reference velocity calculated on velocity data of those frames is
reliable. If such a determination can be done, a reference velocity
is calculated accordingly. Those 14 frames may be consecutive
(refer to (1) and (2) in FIG. 29) or intermittent (refer to (3) in
FIG. 29). Using the calculated reference velocity, moving
velocities detected from a moving element are then corrected for
pulsatile display at pixels subjected to the velocity correction on
each of the frames that have contributed to calculating the
reference velocity.
[0221] In contrast, if velocities of a moving element can be
detected only in frames less than 14 frames among the 20 frames,
the determination can be made such that a reference velocity
calculated on velocity data of those frames is unreliable. If such
a determination can be done, a reference velocity will not be
calculated, thus the velocity correction being not performed (refer
to (4) in FIG. 29). In this case, however, signals are still
detected from blood flowing in a region to be scanned. Therefore,
to show that there is blood flow in the region, the display on the
ordinary power display mode is carried out at pixels in frames that
have been related to the velocity detection. However, neither the
pulsatile display nor the power display is applied to pixels at
which velocities of a moving element have not been detected.
[0222] As a result of it, as shown in FIG. 30(a), both of the
pulsatile display and the power display are conducted by mixture
(hereafter, this display mode is referred to as "pulsatile
flow/power mix display"). This pulsatile flow/power mix display
allows all the detected blood vessels to be depicted, while still
maintaining the display of the pulsatility of blood vessels that
have been subjected to the velocity correction. It is therefore
possible to obtain higher detectability of blood vessels and to
display the pulsatility of the blood vessels with higher
detectability, thus improving a diagnostic performance to a greater
extent.
[0223] In this pulsatile flow/power mix display mode, for enhanced
visibility, it is preferred that a pulsatile flow image and a power
image are both depicted in colors. Color bars for such display are
exemplified in FIGS. 30(b) and 30(c), in which two color bars are
placed together; one is for corrected velocities (i.e., for a
pulsatile flow image) and the other is for a power image.
[0224] The example in FIG. 30(b) shows a combination of the color
bar for corrected velocities, which has been shown in FIG. 10(a),
and a color bar for an ordinary power image. The example in FIG.
30(c) shows a combination of the color bar for corrected
velocities, which has been shown in FIG. 10(b), and a color bar for
an ordinary power image. In each of the two cases shown in FIGS.
30(b) and 30(c), the corrected velocities and power values are
colored differently from each other. Thus, a clear distinction is
provided between a pulsatile flow image and a power image, and a
clear distinctiveness is given to degrees of the pulsation
appearing in a pulsed flow. Both of detectability and a diagnostic
performance will be upgraded largely.
[0225] Furthermore, an upper limit and a lower limit accompany the
color bar for corrected velocities shown in FIG. 30(a) previously
described. This way of attaching the limit values makes it easier
to read out a degree of the pulsation from the color bar.
[0226] In addition, as shown in FIG. 30(d), a setting switch can be
provided on a not-shown operation panel in order to allow an
operator to give desired values to the upper and lower limits for
corrected velocities, whereby the visibility for the pulsation can
be enhanced. As a further modification, an aliasing velocity can be
placed on the image shown in FIG. 30(a), while a setting switch is
installed on a not-shown operation panel as shown in FIG. 30(d).
This configuration makes it possible that, with viewing the image
shown in FIG. 30(a), an operator sets a properly chosen aliasing
velocity, which is able to elevate a depiction performance for the
pulsation.
[0227] As has been described above, the present application
examples are able to not only raise the detectability for blood
flow and its pulsatility but also enhance the visibility for the
pulsation. This is helpful for a greatly improved diagnostic
performance.
[0228] The present invention is not limited to the foregoing
embodiments that are typically shown and various modifications and
alterations may occur to one skilled in the art based on contents
written in the claims, without departing from the spirit of the
present invention. These modifications and alterations pertain to
the claim(s) of the present invention.
* * * * *