JP6231354B2 - Ultrasonic diagnostic apparatus and control method thereof - Google Patents

Description

The present invention transmits an ultrasonic beam into a subject, receives the reflected ultrasonic wave, obtains a received signal, and calculates a blood velocity based on the Doppler transition of the ultrasonic wave based on the received signal. The present invention relates to an ultrasonic diagnostic apparatus that displays a color distribution of blood flow velocity in an observation region and a control method thereof .

  2. Description of the Related Art An ultrasonic diagnostic apparatus that displays an image based on a reception signal obtained by transmitting and receiving ultrasonic waves is known as one of apparatuses that project an image of a subject, particularly a human body. The ultrasonic diagnostic apparatus is usually provided with a function for displaying in color the blood flow distribution in the observation region in the subject based on the received signal.

In this ultrasonic diagnostic apparatus, an ultrasonic probe in which a large number of ultrasonic transducers are arranged that vibrate in response to voltage application and transmit ultrasonic waves and generate voltage signals in response to vibrations caused by ultrasonic waves. A probe with is used. The probe's ultrasonic probe is directed to the body surface of the subject, and each of the ultrasonic transducers constituting the ultrasonic probe is delayed according to a predetermined delay pattern. A burst wave signal composed of a plurality of pulses having a frequency f ₀ is applied. Then, an ultrasonic beam focused on a predetermined depth position with a center frequency f ₀ is transmitted in a predetermined direction from the ultrasonic probe into the subject. The reflected ultrasonic waves are picked up by a plurality of ultrasonic transducers constituting the ultrasonic probe to obtain a plurality of signals, and the plurality of signals are respectively delayed according to a predetermined delay pattern and added to each other. . Thereby, a reception signal as an RF signal representing an ultrasonic beam extending into the subject is obtained. This ultrasonic transmission / reception is repeated a plurality of times, and a phase change Φ (t) due to the Doppler transition of the ultrasonic wave in the meantime is obtained. From the phase change Φ (t) and the center frequency f ₀ ,

Where T is the repetition period of transmission and reception C is the speed of sound V _d (t) is the Doppler velocity (blood flow velocity) in the ultrasonic beam direction
t is the time.
Is calculated. Then, the velocity V _d of each point in the observation area is expressed by red as the blood flow toward the body surface, blue as the blood flow away from the body surface, and the magnitude of V _d as the luminance of the color.

Here, since the center frequency f ₀ of the pulse is constant in the above configuration, when trying to detect the blood flow velocity within a certain depth range in the subject, the repetition frequency PRF of ultrasonic transmission / reception = 1 / T ( The maximum speed Vmax that can be measured is limited according to (T is a repetition period). As the repetition frequency PRF is higher, a faster blood flow can be detected, but the depth of the observation region where the blood flow velocity can be displayed becomes shallower. This is because the next wave cannot be transmitted until the ultrasonic wave is transmitted and the ultrasonic wave is returned and the wave is received.

  Further, in calculating the blood flow velocity, an MTI filter is used to remove components other than the blood flow that are caused by movements of organs and the like, and low frequency components are cut. If the repetition frequency PRF is high, the cutoff frequency of the MTI filter is also high, and the lowest detectable blood flow velocity is high, which may be a problem.

  The Doppler transition due to blood flow is caused by the reflection of ultrasonic waves by a large number of red blood cells, etc. in the blood, but so-called speckle noise is generated by random scattering reflections by a large number of red blood cells, etc., and this causes destructive interference. There are frequent cases where the received signal weakens at the point where it occurs, the calculated phase changes drastically, the variance is large, and the speed cannot be determined accurately.

Japanese Patent Application Laid-Open No. 2004-228620 proposes correcting phase difference by transmitting and receiving ultrasonic waves of two frequencies. However while, in order to calculate the blood flow velocity Vd (t) based on the above equation (1), despite the need to divide the center frequency f _0, how with center frequencies present two It is unknown whether the blood flow velocity is calculated.

JP-A-4-269949

In view of the above circumstances, an object of the present invention is to provide an ultrasonic diagnostic apparatus and a control method therefor that can obtain a blood flow velocity over a wide range with high accuracy compared to the conventional art.

The ultrasonic diagnostic apparatus of the present invention that achieves the above object is
A probe responsible for transmitting an ultrasonic pulse beam into the subject and receiving an ultrasonic wave reflected back within the subject;
A transmitter that transmits a plurality of types of burst wave signals having different pulse transmission frequencies to the probe, and transmits an ultrasonic beam having a center frequency corresponding to the burst wave signal to the probe,
A reception unit that captures ultrasonic reception at the probe and generates a plurality of types of reception signals according to a plurality of types of burst wave signals;
A phase difference is calculated from the complex autocorrelation of each one type of received signal for each one type of received signals generated by the receiving unit, and each one type based on the phase difference. A primary calculation unit that performs primary calculation of blood flow velocity for each received signal, and blood based on a plurality of calculated values of blood flow velocity calculated by the primary calculation unit based on the plurality of types of received signals. a calculation unit that possess the secondary calculation section which performs second-order calculation of flow velocity, and outputs the final blood flow velocity,
A display unit that displays the distribution of the final blood flow velocity output by the calculation unit within the observation region ;
The secondary calculation unit is configured to calculate an overlap range in which two or more calculable speed ranges partially overlap each other among a plurality of calculable speed ranges that can be calculated based on each of a plurality of types of received signals. An average value of the calculated values of blood flow velocity calculated primarily based on each of two or more types of received signals corresponding to two or more calculable velocity ranges is calculated and output as the final blood flow velocity. For a range in which the computable velocity ranges do not overlap with each other, the blood flow velocity that is primarily calculated based on one received signal having the computable velocity range is output as the final blood flow velocity. .

  The ultrasonic diagnostic apparatus of the present invention transmits and receives ultrasonic waves having a plurality of frequencies. For this reason, the range of the blood flow velocity which can be calculated can be expanded. Further, by performing primary calculation of blood flow velocity based on the received signal of each frequency and performing secondary calculation of blood flow velocity based on the calculated values of blood flow velocity of a plurality of frequencies, it is highly resistant to speckle noise. Accurate blood flow velocity is required.

  Here, in the ultrasonic diagnostic apparatus of the present invention, the secondary calculation unit can calculate two or more of a plurality of calculable speed ranges that can be calculated based on each of a plurality of types of received signals. For the overlapping range in which the velocity ranges partially overlap each other, the average value of the blood flow velocity calculated value that is primarily calculated based on each of two or more types of received signals corresponding to the two or more computable velocity ranges Is preferably calculated.

  By calculating an average value of a plurality of calculated values of blood flow velocity, for example, an average value, a median value, etc., a more accurate blood flow velocity resistant to speckle noise is obtained.

  Further, in the ultrasonic diagnostic apparatus of the present invention, the display unit performs calculation based on a received signal obtained by transmitting and receiving ultrasonic waves having the lowest center frequency in addition to displaying the blood flow velocity distribution in the observation region. It is preferable to display a color scale representing the correspondence between blood flow velocity and color, which extends to the highest blood flow velocity possible.

  By displaying up to the maximum blood flow velocity obtained by transmitting and receiving ultrasound with the lowest center frequency, an image with a wide dynamic range of velocity is provided.

  Furthermore, in the ultrasonic diagnostic apparatus according to the present invention, the transmission unit transmits a burst wave signal that causes the probe to transmit an ultrasonic beam focused at different depths in the subject according to the center frequency. Preferably there is.

  The ultrasonic beam has a different depth of focus suitable for narrowing the beam diameter depending on the center frequency. By using this, focusing on different depths in the specimen according to the center frequency can narrow the pulse beam as a whole when combining multiple ultrasonic waves with different center frequencies. However, the S / N can be improved.

  In the ultrasonic diagnostic apparatus according to the present invention, the transmitter may transmit the plurality of types of burst wave signals sequentially to the probe a plurality of times for one type of burst wave signal. Alternatively, the transmission unit may repeat transmission of a plurality of types of burst wave signals to the probe once over a plurality of cycles once for each type of burst wave signal.

As described above, according to the ultrasonic diagnostic apparatus and the control method thereof of the present invention, a highly accurate blood flow velocity is required over a wide velocity range.

It is a block diagram showing the structure of the ultrasonic diagnosing device as an example of the past. It is a schematic diagram of a pulse signal generated by a transmitting unit when transmitting an ultrasonic beam a plurality of times (here, N times). It is the figure which showed an example of the color scale displayed together when the blood flow distribution in an observation area is displayed on a display part. It is the figure which showed an example of the received signal, and the blood flow velocity calculated based on the received signal. It is a block diagram showing the structure of the ultrasonic diagnosing device as one Embodiment of this invention. It is the figure which showed the transmission timing of the pulse signal transmitted from a transmission part. It is the figure which showed the range of the blood flow velocity which can be calculated by two color Doppler processing parts. It is the schematic of an ultrasonic beam when making a focus differ. Is a diagram showing the simulation results of the ultrasound beam shape when employing a single transmission frequency f _0. It is a diagram showing the simulation results of the ultrasound beam shape when employing two transmit frequencies f _1, f _2. Two transmission frequencies f _1, f ₂ and the adoption, and a diagram showing the simulation results of the ultrasonic beam form obtained by forming a separate focus.

  Embodiments of the present invention will be described below.

  Here, first, as a comparative example, a conventional typical ultrasonic diagnostic apparatus will be described, and subsequently, an ultrasonic diagnostic apparatus according to an embodiment of the present invention will be described.

  FIG. 1 is a block diagram showing the configuration of an ultrasonic diagnostic apparatus as an example of the prior art.

  In this ultrasonic diagnostic apparatus 1A, a probe 10, a transmission unit 11, a reception unit 12, and a control unit 13 are provided.

  The probe 10 includes an ultrasonic probe (not shown) in which a large number of ultrasonic transducers are arranged. The ultrasonic probe is addressed to the body surface of the human body as the subject.

The transmission unit 11 generates a pulse signal and transmits the pulse signal toward each of a large number of ultrasonic transducers constituting the ultrasonic probe of the probe 10. This pulse signal is a burst signal having a center frequency f ₀ and a plurality of wavelengths for each pulse signal. The generated burst wave signal is delayed according to the delay pattern determined according to the extending direction of the ultrasonic beam in the human body and the focal depth of the ultrasonic beam, and is applied to each ultrasonic transducer. The Then, an ultrasonic wave is transmitted from each of the multiple ultrasonic transducers, and an ultrasonic beam extending in a desired direction and having a focal point at a desired depth position is generated in the human body by the interference action of the ultrasonic waves. It is sent. The ultrasonic beam sent into the human body is reflected at each depth position in the human body as the ultrasonic beam travels from a shallow position to a deep position in the human body and returns to the ultrasonic probe. It is received by each of the sound wave vibrators. Therefore, in the signal obtained by this reception, the time axis corresponds to the depth in the human body from the ultrasonic probe.

  In the receiving unit 12, the signals obtained by each of the ultrasonic transducers according to the reception by the multiple ultrasonic transducers are also in accordance with the extending direction of the ultrasonic beam, the depth of focus, etc. in the human body. The respective delays are added according to the delay pattern. By doing so, a reception signal as an RF (Radio Frequency) signal representing an ultrasonic beam extending in a desired direction in the human body is generated. The reception signal generated by the reception unit 12 is input to the B-mode processing unit 14 and the MTI filter 15.

  The control unit 13 controls the transmission timing of the burst wave signal from the transmission unit 11 and the reception timing at the reception unit 12. In addition to this, the control unit 13 controls the entire operation of the ultrasonic diagnostic apparatus 1A.

  When displaying a B-mode image, that is, an image based on the ultrasonic reflectance distribution in the human body, on the display unit 19, the transmitter 11 sequentially faces the probe 10 toward the human body in the observation region in the human body. The burst wave signals that are sequentially delayed based on different delay patterns are transmitted to the probe 10 so that different ultrasonic beams are transmitted. Similarly, the reception unit 12 performs delay addition so that an ultrasonic beam in a direction in which the ultrasonic beam is transmitted is generated.

  The reception signal obtained in this way by the reception unit 12 is input to the B-mode processing unit 14, subjected to B-mode image processing, and input to the coordinate conversion unit 17.

  The reception signal obtained by the receiving unit 12 is a signal having a time axis in the depth direction along each ultrasonic beam. On the other hand, the display unit 19 displays an image based on an image signal composed of pixel data arranged in a raster scan direction. Therefore, the coordinate conversion unit 17 converts the signal input from the B-mode processing unit 14 into a signal composed of an array of pixel data suitable for display on the display unit 19.

  The signal output from the coordinate conversion unit 17 is input to the image composition unit 18. This image composition unit 18 composes a B-mode image and a color Doppler image representing blood flow in the human body. A color Doppler image representing blood flow in the human body is usually displayed superimposed on a B-mode image. Only a B-mode image may be displayed without displaying a color Doppler image.

  The image synthesized by the image synthesizing unit 18 is input to the display unit 19, and the image is displayed on a display screen (not shown) for diagnosis.

  In addition, when displaying a color Doppler image representing blood flow in the human body, the following processing is performed in addition to the generation of the B-mode image.

  The ultrasonic wave beam extending in the same direction in the human body from the probe 10 is transmitted a plurality of times (for example, 8 times) and received by the probe 10 so that the delay of the burst wave signal by the transmission unit 11 and the reception unit A delay addition at 12 is performed.

  Transmission / reception in the same direction in the human body is performed for each direction of the ultrasonic beam extending into the observation region.

  The reception signal obtained in this way is input to the MTI filter 15. The MTI filter 15 is a kind of high-pass filter, in which a component caused by Doppler transition of blood flow is passed and a component caused by movement of an organ or the like is cut.

  The received signal after passing through the MTI filter 15 is input to the color Doppler processing unit 16. The color Doppler processing unit 16 calculates a complex autocorrelation of the input received signal, calculates a phase difference from the complex autocorrelation, and calculates a blood flow velocity and the like based on the phase difference. This blood flow velocity or the like is performed for each point on the ultrasonic beam. Since the direction of the ultrasonic beam sequentially changes in the observation region, the blood flow velocity and the like are calculated for each two-dimensional point in the observation region. Data representing the calculated blood flow velocity and the like is input to the coordinate conversion unit 17 and subjected to coordinate conversion, and is superimposed on the B-mode image by the image composition unit 18. Then, the blood flow distribution superimposed on the B-mode image is displayed on the display unit 19. In this blood flow distribution, the blood flow in the direction toward the ultrasonic probe addressed to the body surface of the probe 10 is normally displayed in red, and the blood flow in the direction away from the probe 10 is displayed in blue. Further, the blood flow velocity is displayed with a red or blue color luminance. In this way, the blood flow distribution in the observation area is displayed and used for diagnosis.

  Hereinafter, processing in the color Doppler processing unit 16 will be described.

  FIG. 2 is a schematic diagram of a pulse signal transmission / reception interval generated by the transmission unit when transmitting an ultrasonic beam a plurality of times (here, N times).

Here, N pulse signals including pulse signals 1, 2,..., N-1, N are shown. In this figure, the pulse signals are shown in a simplified manner, but each pulse signal is a burst signal having a center frequency f ₀ and a plurality of wavelengths. Here, the repetition period of the pulse signal is referred to as T, and the repetition frequency is referred to as PRF. There is a relationship of T = 1 / PRF.

  2 is received by the receiver 12 when the pulse signal schematically shown in FIG. 2 is transmitted to the probe 10 with the same delay pattern, and further input to the color Doppler processor 16 via the MTI filter 15. The received signal

However, S (n, t) is an output signal from the MTI filter, n is the number of the transmission / reception signal, and n = 1, 2,..., N−1, N shown in FIG.
t is the time from the reference time for each received signal A (n, t) is the amplitude e ^{jθ (n, t)} , phase j is the imaginary unit I (n, t) is the real part Q (n , T) is the imaginary part.

  In the color Doppler processing unit 16, the complex autocorrelation of the received signal S (n, t)

However, * represents a complex conjugate.
Is calculated. The real part Re and the imaginary part Im of the complex autocorrelation R (1, t) are as follows.

  The color Doppler processing unit 16 further calculates the phase difference from this complex autocorrelation.

Is calculated.

This phase difference Φ (t), Doppler angular frequency ω _d (t), and Doppler frequency f _d (t)

Are in a relationship. Doppler theorem

Where C is the speed of sound.
As can be seen, the Doppler frequency f _d is proportional to the transmission frequency f ₀ . From these relationships, the blood flow velocity, that is, the Doppler velocity in the ultrasonic beam direction, V _d (t), is

Is calculated.

  Further, the color Doppler processing unit 16 can also calculate the power of blood flow Power (t) and the variance Var (t). Blood power Power (t) and variance Var (t) are

Is calculated according to

  FIG. 3 is a diagram illustrating an example of a color scale displayed together with the blood flow distribution in the observation region displayed on the display unit.

V _dmax shown in FIG. 3 is the maximum velocity of blood flowing in an improvement approaching the body surface along the ultrasonic beam among the highest blood flow velocity that can be calculated by the color Doppler processing unit 16. Further, −V _dmax is the maximum velocity of blood flow that flows in the direction away from the body surface along the ultrasonic beam.

  Therefore, this color scale is red in the upper half from the center and blue in the lower half, and becomes brighter as the distance from the center increases. The blood flow velocity at each point in the observation area is displayed in a color and brightness corresponding to the blood flow velocity and direction at that point, corresponding to this color scale.

  Here, the speed range that can be calculated by the color Doppler processing unit 16 will be considered.

  The phase difference Φ (t) is calculated according to the above equation (6). Therefore, this phase difference Φ (t) is

Subject to restrictions. Therefore, the highest Doppler frequency f _dmax is

It becomes. Here, T is a transmission / reception repetition cycle, and PRF is a transmission / reception repetition frequency (see FIG. 2).

  The repetition frequency PRF of this transmission / reception is

Subject to restrictions. Here, Depth is the maximum depth of the observation region, and C is the speed of sound. The ultrasonic wave travels through the human body at the speed of sound C, is reflected, and returns at the speed of sound C. Therefore, when the repetition frequency PRF of transmission / reception is high, it means that the observable depth Depth becomes shallow. From this, the maximum speed V _dmax that can be calculated is

It becomes. This equation (16) means that the maximum speed V _dmax is limited according to the repetition frequency PRF of transmission / reception, that is, according to the observation depth Depth.

Next, the minimum speed V _dmin that can be calculated will be considered.

Minimum speed V _dmin is

It is represented by

Here, fcutoff is a cutoff frequency of the MTI filter 15. Cut-off frequency of the MTI filter 15 is normally Me determined in a percentage K% of the repetition frequency PRF and received. Therefore, Equation (17) is

It becomes.

As described above, when the maximum speed V _dmax and the minimum speed V _dmin are limited according to the transmission / reception repetition frequency PRF and an attempt is made to observe a deep point in the human body, it is necessary to lower the repetition frequency PRF, and high-speed blood flow. Cannot be measured. On the other hand, when the repetition frequency PRF is increased, the minimum velocity V _dmin is increased as can be seen from the equation (18), and low-speed blood flow cannot be observed.

  Next, another problem that the ultrasonic diagnostic apparatus 1A shown in FIG. 1 has will be described.

  The blood flow velocity is calculated based on the Doppler transition when the ultrasound is reflected by a large number of red blood cells in the blood, but speckle noise occurs due to the scattering reflection of the large number of red blood cells, and the received signal is caused by interference. A weakened part occurs. When the received signal weakens, the calculated phase difference fluctuates violently and the blood flow velocity is often not correctly calculated.

  FIG. 4 is a diagram illustrating an example of a received signal and a blood flow velocity calculated based on the received signal. 4A and 4B, the horizontal axis is a time axis, that is, an axis corresponding to the depth of the human body. The vertical axis in FIG. 4A is the amplitude of the received signal, and the vertical axis in FIG. 4B is the blood flow velocity.

  In FIG. 4A, two received signals are shown to be almost overlapped. Here, it is not necessary to distinguish between the two received signals, and an image with a large or small amplitude is sufficient.

  In the received signal shown in FIG. 4A, there are regions a and b in which the received signal is weakened, and the blood flow velocity shown in FIG. 4B is large for the regions a and b, particularly in the region a. It has fluctuated. This is a blood flow velocity that is erroneously calculated due to speckle noise.

  Based on the description of the ultrasonic diagnostic apparatus as a comparative example, the ultrasonic diagnostic apparatus according to an embodiment of the present invention will be described next.

  FIG. 5 is a block diagram showing a configuration of an ultrasonic diagnostic apparatus as an embodiment of the present invention. 5, the same components as those of the ultrasonic diagnostic apparatus 1A as the comparative example shown in FIG. 1 are denoted by the same reference numerals as those shown in FIG. Here, differences from the above-described comparative example will be described.

  The ultrasound diagnostic apparatus 1B shown in FIG. 5 includes two MTI filters 15_1 and 15_2, two color Doppler processing units 16_1 and 16_2, and a combining unit 20. The two color Doppler processing units 16_1 and 16_2 correspond to an example of a primary calculation unit according to the present invention, and the combining unit 20 corresponds to an example of a secondary calculation unit according to the present invention.

  FIG. 6 is a diagram illustrating the transmission timing of the burst wave signal transmitted from the transmission unit.

Here, in order to detect the blood flow velocity up to a certain depth, a necessary PRF is determined and fixed to the PRF. In order to detect a faster blood flow velocity than when using the transmission frequency f ₀ employed in the comparative example (FIG. 1), here, a first frequency f ₁ lower than the transmission frequency f ₀ , here f ₁ = (2/3) and transmits a burst wave signal of _{f 0.} Further, in order to detect a slower blood flow velocity than when the transmission frequency f ₀ is used, a second frequency f ₂ higher than the transmission frequency f ₀ , here, f ₂ = (4/3) f ₀ is transmitted.

FIG. 6A shows one transmission method. Here, after transmission of a burst wave signal of transmission frequency f ₁ is repeated, transmission of a burst wave signal of transmission frequency f ₂ is repeated. .

In FIG. 6B, a burst wave signal having a transmission frequency f _{1 and} a burst wave signal having a transmission frequency f ₂ are alternately and repeatedly transmitted.

FIG. 6 (A), the in any case the transmission scheme of FIG. 6 (B), the received signal when the transmission frequency f ₁ is input to the MTI filter 15 _, filtering a row corresponding to the transmission frequency f ₁ And input to the color Doppler processing unit 16_1. Similarly, the received signal when the transmission frequency f ₂ is input to the MTI filter 15_2 is input to the color Doppler processing unit 16_2 is performed filtering according to the transmission frequency f _2.

In the color Doppler processing unit 16_1, the blood flow velocity V _d 1 is determined from the reception signal of the transmission frequency f _1.

Is calculated by Here, f _d 1 is a Doppler frequency.

  The maximum and minimum blood flow rates calculated by the color Doppler processing unit 16_1 are

However, V _dmax1 and V _dmin1 are the maximum and minimum blood flow velocities for the transmission frequency f ₁ ,
V _{dmax, V dmin} is greatest when the transmission frequency _{f 0,} the minimum blood flow _rate, K 0, _{K 1} is the case of _f 0, _{f 1} transmit frequency, respectively, each define a cut-off frequency of the MTI filters It is a percentage (see equation (18)).

Similarly, in the other color Doppler processing unit 16_2, the blood flow velocity V _d2 is obtained from the received signal of the transmission frequency f ₂ .

Is calculated by However, _{f d} 2 is the Doppler frequency.

  The maximum and minimum blood flow velocity calculated by the color Doppler processing unit 16_2 is

However, V _dmax2 and V _dmin2 are the maximum and minimum blood flow velocities for the transmission frequency f ₂ ,
V _{dmax, V dmin} is greatest when the transmission frequency _{f 0,} the minimum blood flow _rate, K 0, _{K 2} is the case of _f 0, _{f 2} transmission frequency, respectively, each define a cut-off frequency of the MTI filters It is a percentage (see equation (18)).

  FIG. 7 is a diagram showing a range of blood flow velocity that can be calculated by two color Doppler processing units.

FIG. 7A shows a range of blood flow velocity that can be calculated based on the received signal when the transmission frequency f ₁ = (2/3) · f ₀ , and FIG. 7B shows the transmission frequency f _2. The range of blood flow velocity that can be calculated based on the received signal when = (4/3) · f ₀ is shown.

Further, FIG. 7C shows a range of blood flow velocity that can be calculated based on the reception signal at the transmission frequency f ₀ (f ₁ <f ₀ <f ₂ ).

  In the synthesis unit 20 illustrated in FIG. 5, the final blood flow velocity is calculated by the following calculation based on the calculated blood flow velocity values calculated by the two color Doppler processing units 16_1 and 16_2.

When the calculated value of the blood flow velocity is within the range of the blood flow velocity region R1, the calculated values of the color Doppler processing unit 16_1, that is, the blood flow velocity V _d , the power Power, and the variance Var,
V _d = V _d 1
Power = Power1
Var = Var1
However, V _d 1, Power 1, Var 1 are the blood flow velocity, power, and variance calculated from the received signal at the transmission frequency f ₁ .
Is adopted.

In addition, when the calculated value of the blood flow velocity is within the blood flow velocity region R2, the calculated value of the color Doppler processing unit 16_2, that is,
V _d = V _d 2
Power = Power2
Var = Var2
However, V _d 2, Power 2, and Var ₂ are blood flow velocity, power, and variance calculated from the received signal at the transmission frequency f ₂ .
Is adopted.

  Further, when the calculated value of the blood flow velocity is within the blood flow velocity region R0, the calculation can be performed by both of the two color Doppler processing units 16_1 and 16_2. value

Is calculated, and blood flow velocity, power, and dispersion as the average value are adopted.

Thus, in this embodiment, the range in which the blood flow velocity can be calculated is expanded by employing the _two transmission frequencies f ₁ and f ₂ . That is, the maximum and minimum blood flow velocities V _ndmax and V _ndmin that can be calculated by the ultrasonic diagnostic apparatus 1B shown in FIG.

It becomes. However, here, the ratios K ₁ (%) and K ₂ (%) of the repetition frequency PRF that define the cut-off frequency in each of the two MTI files 15_1 and 15_2 are both K ₁ = K ₂ = K ₀ ( (See formula (18)), and the ratio is set to the ratio K ₀ (%) in the ultrasonic diagnostic apparatus 1A of FIG.

  Thus, according to this embodiment, the range of the maximum and minimum blood flow speeds that can be observed can be expanded.

Here, K ₁ = K ₂ = K ₀ , but when it is determined that the clutter component is small as a result of observing the displayed image, K ₁ and K ₂ can be detected so that a slower blood flow can be detected. May be set separately.

Further, the detection accuracy of the blood flow velocity and the resolution of the color Doppler image are in a trade-off relationship. However, in the ultrasonic diagnostic apparatus 1B of the present embodiment, the transmission frequencies f ₁ and f ₂ are switched and used, and therefore, FIG. Compared to the case where a single transmission frequency f ₀ is employed as in the illustrated ultrasonic diagnostic apparatus 1A, the trade-off is alleviated, and the detection speed and resolution are improved.

  Furthermore, in the case of the present embodiment, since ultrasonic beams having a plurality of center frequencies are repeatedly transmitted, the S / N is high, and a large number of Doppler velocities detected by receiving ultrasonic pulses having different center frequencies are used. The detection error of the blood flow velocity due to the scattered reflection from the red blood cells is reduced, and the detection accuracy of the blood flow velocity is improved.

So far, the focal point of the transmitted ultrasonic beam has not been described, but this focal point may be the same for the ultrasonic beam having the transmission frequency f _{1 and} the ultrasonic beam having the transmission frequency f _2. Alternatively, they may be different from each other in the depth direction in the human body.

FIG. 8 is a schematic diagram of an ultrasonic beam when the focus is changed. Here, the focal point of the ultrasonic beam having the transmission frequency f ₁ is formed at a position shallower than the focal point of the ultrasonic beam having the transmission frequency f ₀ . On the other hand, with the ultrasonic beam of the transmission frequency f ₂ is the focal the focal also deep position of the ultrasonic beam of the transmission frequency f ₀ is formed. Note that the entire fan-shaped region D represents the observation region D, and the dotted line L represents the direction in which the ultrasonic beam extends, that is, the traveling direction of the ultrasonic wave.

FIG. 9 is a diagram showing a simulation result of the ultrasonic beam shape when a single transmission frequency f ₀ is employed. Here, the transmission frequency f ₀ = 3 MHz and the focal length = 80 mm are employed. Here, in both FIG. 9A and FIG. 9B, the intensity distribution of the ultrasonic beam is represented by contour lines. 9A is a perspective view, and FIG. 9B is a plan view. The same applies to FIGS. 10 and 11 described later. Further, (a) is an ultrasonic beam on the transmission side, (b) is an ultrasonic beam on the reception side, and (c) is an ultrasonic beam when both transmission and reception are integrated. In any of the drawings, x represents the arrangement direction of the ultrasonic transducers and z represents the depth direction.

FIG. 10 is a diagram showing a simulation result of the ultrasonic beam shape when _two transmission frequencies f ₁ and f ₂ are employed. Here, transmission frequencies f ₁ = 2 MHz and f ₂ = 4 MHz are employed. The focal length is 80 mm when both transmission frequencies f ₁ and f ₂ are used.

Transmitting-side ultrasonic beams shown in FIG. 10 (A) is obtained by superimposing ultrasound beams and the ultrasonic beam of the transmission frequency f ₂ of the transmission frequency f _1. The same applies to the reception-side ultrasonic beam in FIG. 10B and the ultrasonic beam that combines both transmission and reception in FIG.

Comparing FIG. 9 and FIG. 10, in the case of FIG. 10 employing _two transmission frequencies f ₁ and f ₂ , the sound near the focal point is more than in the case of FIG. 9 employing a single transmission frequency f ₀ . It can be seen that the power is strong and the beam width is narrow.

FIG. 11 is a diagram showing the simulation result of the ultrasonic beam shape when two transmission frequencies f ₁ and f ₂ are employed and different focal points are formed. Here, as in the case of FIG. 10, the transmission frequencies f ₁ = 2 MHz and f ₂ = 4 MHz are employed. Regarding the focal point, the focal length = 60 mm for an ultrasonic beam with a transmission frequency f ₁ = 2 MHz, and the focal length = 120 mm for an ultrasonic beam with a transmission frequency f ₂ = 4 MHz.

FIG. 11A shows the transmission side ultrasonic beam having the transmission frequency f _{1 and} the transmission side ultrasonic beam having the transmission frequency f _{2 in} an overlapping manner, as in FIG.

  The same applies to the reception-side ultrasonic beam in FIG. 11B and the ultrasonic beam that combines both transmission and reception in FIG. 11C.

Comparing FIG. 10 and FIG. 11, in the case of FIG. 11, the focal point is changed between the ultrasonic beam having the transmission frequency f _{1 and} the ultrasonic beam having the transmission frequency f ₂ , so that it is within the entire observation region (particularly at a long distance). The sound power is strong and the beam width is narrow. If this is adopted, the resolution of the color Doppler image can be further improved, and the blood flow detection accuracy in the entire region can be improved.

Here, the case where the two transmission frequencies f ₁ and f ₂ are employed has been described. However, for example, the three transmission frequencies f ₁ , f ₀ and f ₂ may be employed, and the present invention may include two or more transmission frequencies. What is necessary is just to employ | adopt a transmission frequency.

DESCRIPTION OF SYMBOLS 1A, 1B Ultrasonic diagnostic apparatus 10 Probe 11 Transmission part 12 Reception part 13 Control part 14 B mode process part 15,15_1,15_2 MTI filter 16,16_1,16_2 Color Doppler process part 17 Coordinate conversion part 18 Image composition part 19 Display part 20 Synthesizer

Claims (6)

A probe responsible for transmitting an ultrasonic beam into the subject and receiving an ultrasonic wave reflected back within the subject;
Transmitting a plurality of types of burst wave signals having different pulse transmission frequencies to the probe, and causing the probe to transmit an ultrasonic beam having a center frequency corresponding to the burst wave signal;
A reception unit that captures ultrasonic reception by the probe and generates a plurality of types of reception signals corresponding to the plurality of types of burst wave signals;
A phase difference is calculated from the complex autocorrelation of each one type of received signal for each one type of received signal generated from the plurality of types of received signals, and each of the received signals is generated based on the phase difference. Based on a primary calculation unit that performs primary calculation of blood flow velocity for each type of received signal, and a plurality of calculated values of blood flow velocity calculated by the primary calculation unit based on the plurality of types of received signals. a calculation unit that possess the secondary calculation section that performs second calculation of blood flow velocity, and outputs the final blood flow velocity Te,
A display unit that displays a distribution in the observation region of the final blood flow velocity output by the calculation unit ;
The secondary calculation unit is configured to detect an overlapping range in which two or more calculable speed ranges partially overlap each other among a plurality of calculable speed ranges that can be calculated based on each of the plurality of types of received signals. The final blood flow velocity is calculated by calculating an average value of the blood flow velocity calculated primarily based on each of the two or more types of received signals corresponding to the two or more computable velocity ranges. As for the range where the computable velocity ranges do not overlap with each other, the blood flow velocity that is primarily calculated based on one received signal having the computable velocity range is outputted as the final blood flow velocity. An ultrasonic diagnostic apparatus. In addition to displaying the blood flow velocity distribution in the observation area, the display unit further extends to the highest blood flow velocity that can be calculated based on the received signal obtained by transmitting and receiving ultrasonic waves with the lowest center frequency. the ultrasonic diagnostic apparatus of placing serial to claim 1, characterized in that to display a color scale that represents the correspondence relationship between blood flow velocity and color. And the transmission unit, the probe, according to claim 1, characterized in that to transmit a burst wave signal to transmit an ultrasonic beam focused at different depths within the subject based on the center frequency or 2. The ultrasonic diagnostic apparatus according to 2. And the transmission unit, to the probe, by a plurality of times in succession for one type of burst wave signal, the sequentially for a plurality of types of burst wave signal, of claims 1 to 3, characterized in that to transmit among ultrasonic diagnostic apparatus according to any one. The transmission unit is configured to repeat the transmission of the plurality of types of burst wave signals to the probe over a plurality of cycles cyclically once for each type of burst wave signal. The ultrasonic diagnostic apparatus according to any one of 1 to 3 . A method for controlling an ultrasonic diagnostic apparatus having a probe for transmitting an ultrasonic beam into a subject and receiving an ultrasonic wave reflected and returned within the subject,
  A transmission step of transmitting a plurality of types of burst wave signals having different pulse transmission frequencies to the probe, and causing the probe to transmit an ultrasonic beam having a center frequency corresponding to the burst wave signal;
  A reception step of capturing a plurality of types of received signals according to the plurality of types of burst wave signals by capturing ultrasonic waves received by the probe;
  For each one type of reception signal generated in the reception step, a phase difference is calculated from the complex autocorrelation of each one type of reception signal, and each of the received signals is calculated based on the phase difference. The primary calculation of the blood flow velocity for each type of received signal is performed, and the blood flow velocity is calculated based on a plurality of calculated values of the blood flow velocity calculated by the primary calculation based on the plurality of types of received signals. A calculation step of performing a secondary calculation and outputting a final blood flow velocity;
  A display step of displaying a distribution in the observation region of the final blood flow velocity output in the calculation step on a display unit,
  The secondary calculation in the calculation step is an overlap in which two or more computable speed ranges partially overlap each other among a plurality of computable speed ranges that can be calculated based on each of the plurality of types of received signals. As for the range, an average value of the calculated values of the blood flow velocity that is primarily calculated based on each of the two or more types of received signals corresponding to the two or more computable velocity ranges is calculated, and the final value is calculated. For a range that is output as a blood flow velocity and the calculable velocity ranges do not overlap each other, the blood flow velocity that is primarily calculated based on one received signal having the calculable velocity range is the final blood flow velocity. A method for controlling an ultrasonic diagnostic apparatus, characterized in that:

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