JPS6096233A - Ultrasonic blood flow measuring apparatus - Google Patents

Abstract

(57) [Summary] This bulletin contains application data before electronic filing, so abstract data is not recorded.

Description

DETAILED DESCRIPTION OF THE INVENTION Field of the Invention The present invention relates to an ultrasonic blood flow measuring device, and more particularly to an ultrasonic surface flow measuring device capable of accurately measuring blood flow inside an inner tube.

BACKGROUND ART A B-mode ultrasound image display device that displays B-mode tomographic images inside a living body in real time using an ultrasound echo method is well known, and it is possible to visually determine the distribution of blood vessels inside a living body from the obtained tomographic images. It is widely used in clinical and other settings today.

On the other hand, an ultrasonic Doppler measuring device that utilizes the Doppler effect of ultrasound is also well known, and this device can measure, for example, the velocity of blood flowing in a blood vessel from the Doppler shift frequency, and is widely used today. There is.

However, although each of these devices is effective in measuring the distribution of blood vessels and blood flow velocity inside one body, they cannot accurately measure the amount of blood flowing inside the blood vessels, and effective countermeasures are desired. It was rare.

In particular, the cross-sectional shape of blood vessels inside a living body is not necessarily constant, and in addition to circular shapes, there are also flat shapes or pathological shapes depending on the type of disease. There has been a problem in that it is not easy to accurately measure blood flow.

Purpose of the Invention The present invention has been made in view of such conventional problems, and its purpose is to monitor blood vessels inside a subject using 8-mode tomographic images, and to determine the amount of blood flowing inside the monitored blood vessels. An object of the present invention is to provide an ultrasonic blood flow measurement device that can accurately measure blood flow regardless of the cross-sectional shape of blood vessels.

Structure of the Invention In order to achieve the above object, the apparatus of the present invention includes a B-mode probe that scans an ultrasonic pulse beam orthogonally to the direction of blood vessel movement within a subject, and B
An image display device that displays an 8-mode tomographic image of a subject from a received mode signal performs complex signal processing on the received B-mode signal and calculates the Doppler of the received B-mode signal from the =3-autocorrelation of the complex signal. a first Doppler measurement device that calculates information; and a calibration probe that is installed at a different position from the B-mode probe and receives an echo beam from the imaged blood vessel tomographic plane as a velocity calibration signal. and a second Doppler measurement device that calculates Doppler information of the received velocity calibration signal, and a first Doppler information calculated by the first Doppler measurement device and a second Doppler information calculated by the second Doppler measurement device. a memory circuit for storing Doppler information No. 2, and a memory circuit that reads Doppler information stored in the memory circuit, determines blood flow velocity from the first and second Doppler information, and includes the Doppler information in the first Doppler information based on the blood flow velocity. The system includes an arithmetic circuit that calibrates the blood flow velocity distribution and calculates the amount of blood flowing in the blood vessel by area-integrating it along the cross section of the blood vessel, and monitors the blood vessel in the subject using 8-mode tomographic images. The method is characterized in that the amount of blood flowing through the monitored blood vessel is measured regardless of the cross-sectional shape of the blood vessel.

Example 4- Next, preferred embodiments of the present invention will be described based on the drawings.

FIG. 1 shows a preferred embodiment of the ultrasonic blood flow measurement device according to the present invention.
It is supplied to the mode probe 220.

The B-mode probe 220 is excited and controlled thereby and transmits the ultrasonic pulse beam 100 toward the subject 240.

The echo beam from the subject 240 obtained by transmitting the ultrasonic pulse beam 100 is converted into an electrical signal by the probe 220, and after a desired amplification effect is performed by the ultrasonic transceiver 200, one of the The output is supplied to the image display device 280 as a B-mode received signal, and the other output is supplied to the first Doppler measuring device 300 for measuring the blood flow flowing in the blood vessel.

In order to perform B-mode image display, the signal supplied to the image display device 280 is transmitted to the CR1 display 360 via the detector 320 and the digital scan converter 340.
and modulates the brightness of the display surface of the CR1 display 360.

In order to enable this beat mode image display, in this embodiment, the control circuit 380 connects the ultrasonic transceiver 20 to the ultrasonic transceiver 20.
A signal for beam scanning control is supplied to the 0 direction, which causes the ultrasonic pulse beam transmitted and received from the B-mode probe 220 to be scanned by mechanical or electrical angular deflection, etc. The subject 240 is periodically scanned with the acoustic pulse beam 100, or scanning is stopped at a desired deflection angle.

Further, from this control circuit 380, a sweep synchronization signal is supplied to the digital scan converter 340 in synchronization with this signal for scanning position control, and sweep control of the CR1 display 360 is performed.

Therefore, as shown in FIG.
By positioning the B-mode probe 220 with respect to the subject 240 so that A before scanning 0 is orthogonal to the blood vessel 260 in the subject 240, the image shown in FIG. As shown in , an 8-mode tomographic image of the subject 240 is displayed. In this way, according to the present invention, it is possible to visually recognize the distribution of blood vessels inside the subject 240 from this B-mode tomographic image.

Further, in the present invention, the other output of the ultrasonic transceiver 200 is subjected to a complex operation to obtain predetermined blood flow information. For this purpose, the B-mode received signal obtained from the ultrasonic transceiver 200 is input to the first Doppler measurement device 300, where it is subjected to predetermined complex signal processing, and the B-mode received signal is determined based on the autocorrelation of the complex signal. The Doppler shift frequency fd of the wave signal is calculated and output as Doppler information.

In this way, in the present invention, since the autocorrelation method is used, Doppler information at each point on the axis of the ultrasound pulse beam 100 is continuously detected from the received signal input via the ultrasound transceiver 200. Is possible.

The first Doppler measurement device 300 using this autocorrelation method has already been disclosed in Japanese Patent Application No. 7-57-070479 previously filed by the applicant of the present application.
FIG. 4 shows a specific example of the first Doppler measuring device 300 using such an autocorrelation method.

That is, in the embodiment shown in FIG. 4, the signal input from the ultrasound transceiver 200 to the first Doppler measuring device 300 is first supplied to the complex signal converter 36 and converted into a complex signal.

In the embodiment, the complex signal converter 36 includes a pair of mixers 38a, 38b including phase detectors, in each mixer 38 the received high frequency signal is operated on as a complex reference signal 102, 104, respectively. Complex reference signal 102.
104 has a frequency that is an integral multiple of the transmission repetition frequency of the ultrasonic pulse beam 100 transmitted and received from the probe 220, and has a complex relationship with different phases. Therefore, the mixer 38 can output a complex signal corresponding to a high frequency signal. That is, each mixer 38 outputs a frequency 8- signal of the sum and difference of image frequencies between the received high-frequency signal inputted by mixed detection and the complex reference signal, and these two signals are passed through the low-pass filters 40a and 4.
0b, and only the difference frequency component is extracted.

In the mixed detection function of the mixer 38, the complex reference signals 102 and 104 are single-frequency continuous waves, but the other input signal, the received high-frequency signal, is a pulse wave containing Doppler information. A large number of spectral components will be present in the output. This complex transformation will be explained below using arithmetic expressions.

One of the multiple search reference signals 102 has a frequency fo that is an integer multiple of the transmission repetition frequency fpr, and when its amplitude is 1, it is represented by a sinusoidal voltage signal of sin (2πfot) - (1), and on the other hand, The received high frequency signal is sin (2yrfo t +2πfd t ) −(
2). However, fd is the Doppler shift frequency.

Note that this received signal has an additional 5 in (2zr(fo±nfpr) t +2πfd(
1±nfpr/fo)t)(1)'),he'yt)
(fpr is the transmission repetition frequency, n is an integer of 011.2...), but to simplify the explanation below, the spectrum shown in equation (2) of the poem where n = Q is I will only explain about.

Since the mixer 38a multiplies the complex reference signal 102 and the received high frequency signal, the following equation is obtained by calculating the product of equations (1) and (2).

cos(2πfdt) −cos(4πfot+
2πfdt) From this output, the high frequency in the second term of the above equation is removed by the low-pass filter 710a, so the output signal becomes cos(2πfdt) (3).

On the other hand, since the complex reference signal 104 has a phase difference of 90° from the signal 102, it is expressed as a cosine wave voltage signal of cos(2πfot) (4), and is expressed by the mixed detection of the mixer 38b and the filtering action of the reduction filter 40b. , sin (2πfd t ) − (5), and is converted into a complex signal with the real part of the equation (3) and the imaginary part of the equation (5), and these two signals are as follows. It can be shown by the complex expression of

11=CO8(2πfdt) 11sin (2πfdt) −(6) The signal 11 complex-converted as above is sent to the AD converters 42a, 4
2b, it is converted into a digital signal and output to the next stage complex delay line canceller 44. The AD converter 42/\ is supplied with a clock signal 108 and performs sampling using the clock signal.

In this embodiment, since the above-described complex delay line canceller 44 is provided, it is possible to remove the received signal from the stationary part or the low-speed moving part and extract the speed signal of only the moving part, and the signal in image 11- Quality can be significantly improved. That is, in general, for example, blood flow signals from a living body are mixed with reflected signals (clutter) from almost stationary living tissues such as blood vessel walls and heart walls, and this signal is weaker than reflected signals from blood flow. It is usually strong and causes significant interference with self-current measurements. However, in this embodiment, such low speed signals can be removed by the complex delay line canceller 44, so that it is possible to detect only signals from the moving part.

The complex delay line canceller 44 has delay lines 46a and 46b having a delay time that corresponds to one cycle (T) of the ultrasonic transmission repetition signal, and the delay line has a delay time that corresponds to the number of clock pulses included in one cycle, for example. can be formed from a memory or shift register consisting of storage elements equal to . Difference calculators 48a and 48b are connected to these delay lines 46, respectively, and the difference calculators 48 successively compare the input, ie, the current signal, of the delay line 46 with the output, ie, the signal from 12-1 period before. 1 of the signals that fall at the same depth
Calculate the difference between periods. Therefore, since there is no change or a small change between the current signal and the signal one cycle ago in the reflected signal from stationary or low-velocity living tissue, the difference output of the difference calculator 48 is close to zero; For example, a differential output of a blood flow signal having a high velocity is detected as a large value, and thereby the above-mentioned clutter can be reliably suppressed.

The operation of the complex delay line canceller 44 will be explained below using an arithmetic expression. In the figure, the input to the multi-route delay line canceller 44 is a digital signal, but in order to simplify the explanation, the analog signal of equation (6) will be used for explanation of the calculation formula.

When the input Z1 of the delay line 46 is expressed by equation (6), 1
The period-delayed output is Z2 = CO8 (2πfd (t - T) ) + 1 sin
(2πfd (t −T) ) =−(7), and as a result, the difference calculation output of the difference calculator 48 is Z 3 =Z + −22=−2sin (27rfd
T/2)-sin (2πfd (t +T/2))
10i 2sin (2πfd T/2)・cos(2π
fd (t+T/2>), and if the differential output Z3 is expressed as z3-x3+iy3, each of X3 and V3 becomes the following equation.

X 3 =-2sin (2πfd T/2)-sin
(2yrfd (t + T/2) )−(8)y 3
-') sin (2πfd/2) cos(2πrd
(t +T/2>)... (9) As shown in -V, the output of each difference calculator 48a, 48b is
, v3 are output.

The complex signal from which the low-speed signal has been removed as described above is
Next, an autocorrelator 50 performs arithmetic processing to determine the autocorrelation of Z3 with the amount of delay as the king.

First, the input signal z3 is delayed by one period by the delay line 52a152b, and an output 14 is obtained. This output cuff 4 is expressed by the following formula.

z4=x4 Do 1Va x 4=-2sin (2πfd T/2>-5in
(2yrfd (t-T/2) )...(10) V4 =2sin (2πfd T/2)・cos(2
πfd (t-T/2))...(11) And 14
*If -X4-IV4, the correlation can be found by the following formula.

Z3Z4 ”=(X3 +1V3)(X4-1V4)=
X3X4 +VsV4+i (X4V3-X3Va)
In order to find this correlation, the autocorrelator 50 includes 4
Multipliers B54a, 54b, 56a156b and adders/subtractors 58a, 58b are provided to perform the correlation operation.

If the output of the adder/subtractor 58a is R, the above (8) and (9
), (10), and (11), and if the output of the adder/subtractor 58b is I, 1=X4V3−X3V4=4sin 2 (2yrfd T/2 )−sin (
2πfd T) (13) is obtained, and by combining the outputs of both adders and subtracters 58, the correlation is expressed by the following equation.

S=R+iI (14) Since this correlation output S includes signal fluctuation components and noise components generated from the equipment, an averaging circuit averages the correlations to remove these noise components, and this correlation The average is expressed as s=p+*r.

The average circuit adds the output delayed by one period to the current input in the delay lines 60a and 60b in adders 62a and 62b, and repeats the operation of supplying this output to the delay line 60 again. By outputting the bits, we can obtain the average value. In order to improve the accuracy of the average by repeated operations at this time, in the embodiment, weighting circuits 64a and 64b are used.
is provided to attenuate the output and add it to the input. In other words, if the attenuation amount is α, then the current signal is reduced by, for example, 1
The signal from 0 cycles ago is subtracted by αIll and added as the current signal, so the influence on the output is small.
It becomes possible to perform an averaging function similar to that of a reduction filter or a moving average circuit. Furthermore, by changing the weighting amount of the weighting circuit 64, it is possible to change the degree of averaging.

As described above, in this embodiment, the average correlation is obtained from the autocorrelator 50, and this correlation output is sent to the argument calculator 66.
The argument angle θ of the correlation average output S can be determined by . That is, the declination angle θ is determined from equations (12) and (13) as θ-tan-'(D/R)-2yrtaT-(15), and this unity and Doppler shift frequency axis is expressed as θ-tan-'(D/R)-2yrtaT-(15). ... (16) It can be very easily determined from the above-mentioned declination angle θ. That is, since the transmission repetition period T is a constant, the argument angle θ
is proportional to the Doppler shift frequency column, that is, the blood flow velocity, and since the correlations I and R take positive and negative values, respectively, the deviation angle θ can be measured by the amount ±π, which allows us to determine the direction of motion. It is possible to reduce sex by 1q.

The deflection angle θ in the present invention is −r based on equation (15).
-1R, by creating a table in which the values of the argument angle θ corresponding to the possible values of T and H are written in the ROM in advance, and reading out the argument angle θ corresponding to the input page and page from this table. It is possible to perform high-speed calculations.

The deflection angle θ obtained as described above is converted into a Doppler shift frequency fd by the converter 68 and output.

Further, in the above embodiment, the autocorrelator 50 averages the correlations, but it is also possible to directly calculate the deviation angle from the correlations before averaging, thereby obtaining the Doppler shift frequency fd. It is.

As explained above, according to the first Doppler measuring device 300 using the autocorrelation method, by transmitting and receiving the ultrasonic pulse beam 100, the movement of the biological moving part located in the passage line of the pulse beam 100 is Since the velocity distribution is continuously observed, for example, if a blood vessel 260 is present on the passing line of the pulsed beam 100 as shown in FIG. 3(A),
It is possible to continuously measure the velocity distribution of the blood flow flowing through the blood vessel 260 as a Doppler shift frequency.

FIG. 3(B) shows an example of the blood flow velocity distribution determined in this way, and in the figure, the X-axis direction is the direction of the subject 240.
It represents the depth of the echo beam from inside, and the X-axis direction represents the magnitude of the Doppler shift frequency.

As is clear from the figure, the ultrasonic pulse beam 100
It is understood that when the ultrasonic pulse beam 100 is transmitted toward the blood vessel 260 of the subject 240, the blood flow velocity distribution within the blood vessel 260 on the passage line of the ultrasonic pulse beam 100 is continuously observed as the Doppler shift frequency. 19- I can do it.

In particular, in the present invention, since the ultrasound beam 100 is orthogonally scanned from the B-mode probe 220 to the blood vessel 260 in the subject 240, as shown in FIG.
Blood flow velocity distribution at all points on the blood vessel tomographic image plane displayed on the T display 360 can be measured.

The blood flow velocity distribution thus measured by the first Doppler measuring device 300 is output to the memory circuit 400 as first Doppler information.

In the embodiment, this memory circuit 400 has a first memory 420 and a second memory 440, and the memory circuit 400 has a first memory 420 and a second memory 440.
Doppler information is sequentially written and stored in the first memory 420 corresponding to each beam. The flow velocity distribution will be written and stored.

Therefore, by reading out the blood flow velocity distribution included in the Doppler information stored in the first memory 420 and calculating the area along the cross-sectional area of the blood vessel 260, the inside of the blood vessel can be measured. It becomes possible to reduce the flow of blood in ml.

However, in the present invention, the B-mode probe 220
As shown in FIG. 2, the ultrasonic pulse beam 100 is orthogonally scanned with respect to the blood vessel 260, so the blood flow velocity distribution included in the Doppler information stored in the first memory 420 is based on the blood flow It only represents relative speed, not absolute speed. For this reason, as mentioned above, even if we calculate the amount of blood flowing through the private fault plane, this only represents the relative amount of blood flow, and in order to calculate the true amount of blood flow, we need to
It is necessary to determine the absolute blood flow velocity surface within 60 by some means and calibrate the relative blood flow rate calculated above based on the absolute blood flow velocity.

Therefore, in the present invention, the B-mode probe 2
A calibration probe 460 is provided at a position different from that of the echo beam 10 from the tomographic plane of the blood vessel 260 displayed as an image.
0b is received by the probe 460 as a speed calibration signal.

In the embodiment, this calibration probe 460 is perpendicular to the beam scanning plane A of the B-7 probe 220 and
60 is located in a plane including the center of the cross section, and the calibration probe 460 receives waves of
The configuration is such that the angle θ between the echo beam 100b and the beam scanning plane A of the B-mode probe P220 can be determined from the reception timing of 0b.

The T low beam 100b received by the calibration probe 460 is converted into a predetermined electrical signal, and after being amplified in a predetermined manner by the ultrasonic receiver 480, the T low beam 100b is sent to the second Doppler measuring device. 500 is entered. This second Doppler measurement device M500 has almost the same configuration as the first Doppler measurement device 300, and performs complex signal processing on the input signal, and uses the autocorrelation of the complex signal to obtain the result shown in FIG. 3(C). As shown, the second Doppler information represented by the Doppler transition frequency is outputted to the memory circuit 400.

The memory circuit 400 writes and stores the second Doppler information thus input into the second memory 440.

Here, generally the central portion 260- within the blood vessel 260 exhibits the maximum blood flow velocity, and therefore, as shown in FIG. 3(B).
As shown in FIG. 3(C), the first Doppler information 100a shows the maximum Doppler shift frequency Δf I+ at the position t I+ corresponding to the central portion of the blood vessel 260-, and similarly as shown in FIG. The Doppler information 100b of No. 2 indicates the central part of the blood vessel 260-
The maximum Doppler shift frequency Δ at the position t 2+ corresponding to
Indicates f 2+.

Therefore, if the maximum Doppler shift frequency Δf 11 is read out from the first Doppler information stored in the memory 400 and the maximum Doppler shift frequency Δf2+ is read out from the second Doppler information 100b, the Doppler shift frequency Δf
I+ and Δf2+ will represent the Doppler shift frequency from the maximum blood velocity position obtained by different probes 220 and 460, respectively.

Therefore, the arithmetic circuit 520 of the present invention reads these Doppler shift frequencies Δf I+ and 23−Δf'2+ from the memory 400. Come here, blood vessel 260
The calculation of the maximum velocity of blood flowing through the body has already been disclosed in Japanese Patent Application No. 57-029833 previously filed by the applicant of the present application, and its value can be determined by the following equation.

y ,, -(C/fo sinθ) (Δfu'(1+cosθ)+2Δf212-2Δf
IIΔf 2+ (1+cosθ))4 where,
The angle θ of the echo beams 100a and 100b received by the probes 220 and 460, respectively, is determined by the following formula from the maximum shift frequency positions t11 and t2+ detected within each Doppler information 100a and 100b. Detected based on

COSθ=t n /121 Then, the calculation circuit 520 converts the blood flow velocity distribution included in the first Doppler information 100a from the relative velocity distribution to the absolute velocity distribution based on the maximum surface flow velocity V I+ determined in this way. Calibrate to That is, the blood flow velocity distribution included in the first Doppler information 100a is proportional to the absolute velocity distribution of blood flow, as shown in FIG. 3(B). Therefore, the first Doppler information 100a
The first Doppler information 10 is determined by the ratio of the maximum blood flow velocity contained in
By calibrating 0a, it is possible to obtain the absolute velocity distribution of the blood flow velocity distribution included in the first Doppler information 100a.

After the arithmetic circuit 520 of the present invention calibrates the blood flow velocity distribution included in the Doppler information 100a in this way,
By area-integrating this along the blood vessel cross section, blood vessel 2
The absolute value of the blood flow flowing through 60 has been calculated.

That is, the CRT display 3 is stored in the first memory 420.
The first image covering the entire B-mode tomographic image displayed on 60
Therefore, the Doppler information 100a of the blood vessel 260 is stored, and therefore, the blood flow velocity distribution information flowing within the cross section of the blood vessel 260 is read out from this first mesh 420, and while calibrating this to the absolute velocity distribution, the information is read out along the cross section of the blood vessel 260. By integrating the area,
It becomes possible to calculate the absolute value of the blood flow flowing within the cross section of the blood vessel 260.

Thus, according to the present invention, the first Doppler information 100
In order to calculate the amount of blood flowing through the blood vessel 260 by calculating the area of the blood flow velocity distribution contained in the blood vessel 260 in the cross section of the blood vessel 260, the blood flowing through the blood vessel 260 regardless of the cross-sectional shape of the blood vessel 260 is calculated. It becomes possible to accurately measure the flow rate. Then, the calculation circuit 520 displays the blood flow rate calculated in this manner on the CRT display 360 digitally or graphically.

In particular, according to the present invention, it is possible to accurately measure the blood flow flowing through a blood vessel with a flat cross-sectional shape, where it has been difficult to accurately measure the blood flow, or the blood flow at an abnormal location of the blood vessel, such as the position of an aneurysm. becomes possible.

Note that in order to determine the absolute value VI+ of the blood flow velocity described above, it is sufficient to detect the Doppler shift frequency at approximately the center of the cross section 260- within the blood vessel 260. Therefore, the second Doppler measurement device 500 is not limited to the device using the autocorrelation method of this embodiment, but can also be configured to measure only the Doppler shift frequency Δf2+ at the central portion 260- of the cross section of the blood vessel 260. Good too.

Further, in the present invention, the case where one blood vessel 260 to be measured is displayed on the CR1 display 360 has been described as an example, but when actually using the blood flow device of the present invention, Cross-sections of a plurality of blood vessels are often displayed on the CRT display 360. In such a case, the first and second Doppler information may be collected from the blood vessel 260 by marking the cross section of the desired blood vessel displayed on the screen using a marker or the like, for example.

Doppler information collection using such markers is generally performed in conjunction with a microcomputer, and the calibration probe 460
is controlled to receive the echo beam 100b from the marked blood vessel cross section, and the arithmetic circuit 520 is controlled to read out first Doppler information corresponding to the marked blood vessel cross section position from the first memory 420. .

Therefore, the calibration probe 460 used in the present invention
It is desirable to use one that has a sufficiently wide directivity, particularly in the direction of the running force 27 of the blood vessel 260.

Further, as the probe 220 used in this embodiment, it is possible to use a linear scanning type, a sector scanning type, or any other type of probe.

Furthermore, in the above embodiments, a reception-only probe was used as the calibration probe 460, but the present invention is not limited to this. ,
It is also possible to use one that is capable of both transmitting and receiving ultrasonic pulse beams.

FIG. 5 shows another embodiment of the present invention, in which the calibration probe 460 of the embodiment is perpendicular to the ultrasonic scanning plane A of the B-mode probe 220. The ultrasonic pulse beam is formed to scan sectors along the line B. The B-mode received signals received by both probes 220 and 460 are displayed as B-mode tomographic images on the CRT display 360, respectively.

Here, as shown in FIG. 5, the B-mode probe 220 is positioned so that the ultrasonic pulse beam 100 is scanned orthogonally to the blood vessel running direction, and the calibration probe 4
60 is positioned so that its ultrasonic pulse beam is scanned along the running direction of the blood vessel 260, the CRT display 36
0, as shown in FIG. 6(A), an orthogonal tomographic plane of the blood vessel 260 is displayed as a B-mode tomographic image, and at the same time, as shown in FIG. 6(B), a lateral tomographic plane of the blood vessel 260 is displayed. will be displayed as a B-mode tomographic image.

In this way, according to the device of this embodiment, the CRT display 3
60 shows the tomographic plane of the blood vessel 260 to be measured as shown in FIG.
As shown in A), as an orthogonal tomographic plane, as well as in Fig. 6 (
As shown in B), the image is displayed as a lateral tomographic plane, making it possible to visually and multifacetedly recognize information about the blood vessel 260 that is the object of measurement.

In the apparatus of this embodiment, a microcomputer or the like is used to extract the 8-mode reception signals obtained through the probes 220 and 460 as shown in FIG.
A position 260- of the center of the blood vessel tomographic plane shown in A) is automatically specified. At the same time, 8
The echo beams received by the mode probe 220 and the calibration probe 460 from the point 260- are
Images are displayed as beam lines 100a and 100b, and flow rate measurement points are displayed.

Then, the amount of blood flowing through the blood vessel 260 at the measurement point 260- is calculated in the same manner as in the above embodiment,
This will be displayed as an image on the CRT display 360.

In this embodiment, a sector scanning type probe is used as the probe 220.460, but the probe is not limited to this, and a linear scanning type probe may be used.

As described in detail, according to the present invention, blood vessels within a subject can be monitored using B-mode tomographic images, and the amount of blood flowing through the monitored blood vessels can be accurately measured. becomes.

In particular, according to the present invention, an ultrasonic pulse beam is orthogonally scanned over a blood vessel within a subject, and the blood flow velocity distribution included in the obtained Door Arara information is area-integrated along the cross section of the blood vessel. Because it calculates the amount of blood flowing through the
It becomes possible to accurately measure the amount of blood flowing through the blood vessel regardless of the cross-sectional shape of the blood vessel.

[Brief explanation of the drawing]

Fig. 1 is a block diagram showing a preferred embodiment of the ultrasonic blood flow measurement device of the present invention, Fig. 2 is an explanatory diagram of ultrasonic pulse beam scanning by the device shown in Fig. 1, and Fig. 3 (A) is An explanatory diagram of a B-mode tomographic image of the apparatus shown in FIG. 1; FIGS. 3(B) and (C) are explanatory diagrams of waveforms of Doppler information obtained via a B-mode probe and a calibration probe; FIG. 4 is a block diagram showing the detailed configuration of the first Doppler measurement device of the device shown in FIG. 1, FIG. 5 is an explanatory diagram showing another embodiment of the present invention, and FIG. FIG. 3 is an explanatory diagram of a B mode 31-1 dot tomographic image obtained by the apparatus shown in FIG. 220... B-mode probe 240... Subject 260... Blood vessel 280... Image display device 300... First Doppler measurement device 400...
Memory circuit 460... Calibration probe 500... Second Doppler measurement device 520...
Arithmetic circuit A: Beam scanning surface of the B-mode probe. Applicant Aloka Co., Ltd. 32- ~ , * E1t ■ Sniff 禦 81 1″″1ゞ区呆 2°9 ″ ~ Dimension 1 ooo 'Lm'' Ig, *Eco1 1 :: To `` '111 ・1 ■ -〇 na-8ya Tsuba: vessel, dormitory rot 1srQ Shimuga 1 oo ku 101 - 'R+''!・08 Bag "P" - Figure 2 Figure 3 (A) (B) (c)

Claims (1)

[Scope of Claims] (1) A B-mode probe that scans an ultrasonic pulse beam orthogonally to the direction of blood vessel travel within the subject, and a B-mode received signal obtained through this probe. an image display device that displays a B-mode tomographic image of a specimen; a first Doppler measurement device that performs complex signal processing on the B-mode received signal and calculates Doppler information of the B-mode received signal from the autocorrelation of the complex signal; ,
A calibration probe that is installed at a different position from the B-mode probe and receives an echo beam from the imaged blood vessel tomographic plane as a velocity calibration signal, and Doppler information of the received velocity calibration signal. a second Doppler measurement device that calculates; a memory circuit that stores the first Doppler information calculated by the first Doppler measurement device and the second Doppler information encoded by the second Doppler measurement device; The Doppler information stored in the memory circuit is read out, the blood flow velocity is determined from the first and second Doppler information, the free flow velocity distribution included in the first Doppler information is calibrated based on the blood flow velocity, and this is adjusted to the curved tube. an arithmetic circuit that calculates the age of blood flow flowing in a blood vessel by integrating the surface area along a cross section; An ultrasonic blood flow measurement device that measures blood flow regardless of the cross-sectional shape of blood vessels. (2. In the apparatus described in claim (1), the calibration probe is located in a plane that is perpendicular to the beam scanning plane of the B-mode probe and includes blood vessels within the subject. (3) In the device according to any one of claims (1) and (2), the calibration probe is placed on the ultrasonic scanning surface of the B-mode probe. On the other hand, an ultrasonic blood flow measuring device characterized by orthogonal scanning of an ultrasonic beam.