WO2001026555A1 - Ultrasonic imaging device - Google Patents

Abstract

An ultrasonic imaging device having refraction correction delay data creating means for creating a delay time for focusing the transmitted or received wave in consideration of the effect of refraction of ultrasonic wave by an ultrasonic propagation medium between arrayed vibrators and the preset focal point and supplying the delay time to delay control means so as to create an excellent ultrasonic image while preventing any focal shift of the ultrasonic wave transmission/reception due to refraction of the ultrasonic wave caused at the interface between a lens layer of an ultrasonic probe and the fat layer of the subject and at the interface between the fat layer and the biological tissue present under the fat layer.

Description

 Description Ultrasonic imaging equipment Technical field

 The present invention relates to an ultrasonic imaging apparatus for extracting a tissue in a subject as an image using ultrasonic waves, and particularly to an ultrasonic beamformer used for the same. Background art

 An ultrasonic device, for example, an ultrasonic imaging device used for medical image diagnosis obtains a tomographic image of a soft tissue of a living body or an image of a blood flow flowing through the living body in almost real time using an ultrasonic pulse reflection method. It can be displayed on a monitor and observed, and it is said to be highly safe because it does not expose the subject to radiation, unlike a diagnostic imaging device that uses radiation. Applied in the field. In an ultrasonic imaging apparatus, an ultrasonic probe is used for transmitting an ultrasonic wave into a subject and receiving an echo signal from inside the subject. One of the scanning methods of the ultrasonic imaging apparatus is an electronic scanning method. The electronic scanning type ultrasonic probe arranges elongated rod-shaped transducers in a one-dimensional array, and gives a predetermined delay time to each transducer for driving. As a result, the ultrasonic probe transmits an ultrasonic beam converging at a predetermined depth and a predetermined direction in the subject. The received wave is obtained by synthesizing a received signal from each transducer of the ultrasonic probe with a given delay time for each transducer and combining the received signals from a given depth and direction. The processing part that sets the delay time for each transducer and gives the delay time to each transducer is called an ultrasonic beamformer.

In order to obtain a good ultrasonic image by electronic scanning, it is necessary to provide an accurate delay time for each transducer and transmit it over the entire scanning range of the ultrasonic beam. Only with this method is it possible to focus the ultrasound at the intended focal point without distorting or spreading it. The same operation is performed on the received wave to form the beam of the received echo signal. The ultrasonic probe uses ultrasonic waves generated by many transducers arranged in a one-dimensional array. A lens layer for converging in a direction orthogonal to the arrangement direction is provided. In particular, since the phasing circuit of the ultrasonic imaging apparatus has recently been digitized, it has become possible to easily and accurately control the delay time and phasing in the transmission and reception of ultrasonic waves. However, the ultrasonic waves transmitted from each transducer do not converge at the focal point only through a medium with a constant sound velocity. The transmitted pulse is converted from an electric signal into an ultrasonic wave by the piezoelectric vibrator, and reaches a desired focal position through the lens layer on the probe surface and the living tissue. At this time, since the sound speed is different between the lens layer and the living tissue, sound refraction occurs at the interface between the two in the arrangement direction of the piezoelectric vibrators. In calculating the delay time, it is necessary to incorporate the effect of this refraction. Since the refraction angle is obtained by Snell's law, it is possible to calculate the delay time by obtaining the refraction path of the sound.

 The ultrasonic imaging device captures tomographic images of about 30 frames per second while moving the focus position during the capturing of one frame by 50 in the horizontal direction (azimuth direction) and in the depth direction (distance direction). Needs to be changed by about 20. The number of transducers in a single aperture varies from 32 to 192, depending on the transducer. From these facts, it is necessary to set the delay time for one vibrator at a very high speed even if it is calculated. However, there is a limit to the amount of computation that can be performed in real-time with the MPU (Micro Processor Unit) mounted on a real ultrasonic imaging device, even in light of the arithmetic processing performance of a DSP (Digital Signal Processor). For this reason, in the current ultrasonic imaging apparatus, a method of storing data in which the delay time for each focus position is calculated in advance for each probe in a storage medium such as a hard disk, or a method for real-time processing. One of the methods is to perform an approximate calculation using a formula.

At present, probes that are specialized for the organ, site, and symptom to be diagnosed are used as probes of the ultrasonic imaging apparatus, and the types thereof are increasing. Therefore, a considerably large-capacity storage device is required to store the delay time data. However, it is not realistic to mount an extremely large-capacity storage device on an ultrasonic imaging device due to price and other restrictions. In particular, difficulties arise when the ultrasonic imaging device has expandability such that a probe developed later can be used. Also, in the two-dimensional probe, which will become the mainstream in the future, the number of transducers will increase further. Data amount is increasing.

 On the other hand, the method of obtaining the delay time by approximation calculation may be insufficient in accuracy with the conventional calculation accuracy depending on conditions. Since the accuracy is determined by the ratio between the period at the center frequency of the probe and the error time, a probe that transmits a pulse with a high center frequency, that is, a pulse with a short period, like a high-definition probe has been developed. The demands on the accuracy of the delay time become stricter. Therefore, a more accurate approximation method is required. In addition, considering the application to a two-dimensional probe, it is necessary to shorten the calculation time for one transducer, so going in a more severe direction is the same as the former case.

 The image used for ultrasonic diagnosis has been significantly improved in comparison with the conventional image in combination with the adoption of the digital phasing technology, but the image obtained by the ultrasonic imaging apparatus is an X-ray apparatus. There is a demand for further improvement in image quality when compared with images obtained by other modalities such as X-ray CT and MRI.

 As a measure for improving the image quality of the ultrasonic image, a technique for correcting the sound speed of the ultrasonic wave in the subject has been recently developed. In a conventional ultrasonic imaging apparatus, delay control data for focusing an ultrasonic wave to a focal position in a subject is set based on an average sound velocity of a living body. Due to individual differences such as those with a lot of fat, when the control data is used for all subjects, an image that is good on average is obtained, but the best image is obtained for each subject. It could not be said that it could be improved, and its improvement was desired. The above-mentioned sound velocity correction technology is a countermeasure for this. By measuring or estimating the ultrasonic wave propagation velocity in the body for each subject to be examined, the delay time control data for beam formation is calculated based on the measured or estimated value. Is what you want. This sound velocity correction technique is described in, for example, Japanese Patent Application Laid-Open Nos. 8-317923 and 10-66694 filed by the present applicant.

However, it is considered that there is still room for improvement in the image quality of ultrasonic images even if this sound velocity correction technology is adopted. The inventors of the present invention believe that this can be improved by forming a beam by incorporating the refraction in the ultrasonic wave propagation path. As the refraction in the ultrasonic wave propagation path, the refraction by the lens layer is the first problem, but the refraction of the ultrasonic wave occurs only at the interface between the lens layer and the subject. Les ,. The living body to be diagnosed by the ultrasonic imaging device is composed of various tissues such as fat, muscle fibers, various organs, and blood. Since the speed of sound is different in each of these tissues, subtle refraction of ultrasonic waves occurs at the interface. Of these refractions, it is easiest to incorporate the refraction effect of the fatty layer into beamforming.

 There are three main reasons: First, since the fat layer always exists under the skin, the ultrasonic pulse penetrates the fat layer when the ultrasound probe is applied percutaneously to see any of the skin. In other words, refraction by the fat layer cannot be avoided unless a transcutaneous probe is used. Second, since the fat layer is the outermost part of the living tissue, the effect of refraction on the focal position has the greatest effect. Thirdly, it is possible to assume that the subcutaneous fat has a certain thickness within the caliber of the probe, so it is easier to incorporate it into the calculation than in the case of other tissues such as blood vessels. is there. Disclosure of the invention

 The present invention has been made in view of the above, and a first object of the present invention is to further improve the image quality of an ultrasonic image compared to the current state.

 A second object of the present invention is to provide an ultrasonic imaging apparatus capable of setting a delay time in consideration of the effect of ultrasonic refraction by a lens layer and / or a fat layer of an ultrasonic probe. It is in.

 Furthermore, a third object of the present invention is to set a delay time in consideration of the influence of refraction by the lens layer or / and fat layer and the ultrasonic wave propagation speed of each subject. An object of the present invention is to provide an ultrasonic imaging apparatus.

When considering the refraction of ultrasonic waves by the fat layer, if it is necessary to calculate the delay time in advance and store it in the storage device, it is necessary to calculate all for each fat layer thickness. A large-capacity storage device is required. If the sound velocity of the fat layer is slightly different due to individual differences, it is advantageous to be able to calculate on the spot. Therefore, the present invention employs a method of obtaining a delay time by using an algorithm having a sufficient approximation accuracy and a fast calculation, and taking into account the influence of ultrasonic refraction by the lens layer and the fat layer. That is, an ultrasonic imaging apparatus according to a representative example of the present invention includes: an ultrasonic probe having an arrayed transducer; and transmitting, focusing, and receiving when transmitting or Z and receiving ultrasonic waves to a subject. Delay control means for controlling a delay time with respect to each transducer for performing focusing; and the transmitting or transmitting by incorporating a refraction effect of ultrasound by an ultrasound propagation medium between the arrayed transducers and a set focal position. It is characterized by comprising a refraction correction delay data generating means for generating a delay time for performing forcing of a received wave and supplying the delay time to the delay control means, and a display unit for displaying an ultrasonic image. The delay control means stores in advance delay time data obtained based on the average sound velocity of the living body, and performs ultrasonic transmission and reception for obtaining refraction correction data using the stored delay time data in advance. Will be The refraction correction delay data generating means uses parameters relating to the ultrasonic probe including a lens layer thickness of the probe, a sound velocity of the lens layer, and an arrangement pitch of the transducer, and calculates a relationship between the transducer and a designated focal point. The delay time to be given to each transducer is calculated by calculation, taking into account the ultrasonic refraction effect in the ultrasonic propagation path between them. Further, the refraction correction delay data generating means includes: a parameter relating to the ultrasonic probe including a lens layer thickness, a lens layer sound velocity, a pitch between transducers; a fat layer thickness of the subject and a sound velocity of the fat layer; The delay time to be given to each transducer is calculated by using the sound velocity data of the weave, taking into account the ultrasonic refraction effect in the ultrasonic propagation path between the transducer and the designated focal point. Further, the refraction correction delay data generation control means may calculate the delay time using a parameter recursively solved from a parameter relating to a sound path from the vibrator next to the vibrator to be calculated to the focal point. Obtained by calculation.

 Further, the ultrasonic imaging apparatus of the present invention comprises: means for detecting data relating to a layer structure of a subject for refraction correction from a screen of the display unit on which an ultrasonic image is displayed; and an output of the layer structure data detecting means. Means for generating delay control data taking into account the effect of ultrasound refraction due to the layer structure of the subject using the method. The layer structure data detecting means includes a caliper for displaying two movable cursors on the screen and measuring a distance between the cursors on the screen.

And the ultrasonic imaging apparatus of the present invention comprises: an ultrasonic probe having an arrayed transducer; Delay control means for controlling a delay time for each transducer to perform transmission focusing or reception focusing at the time of transmitting or z and receiving ultrasonic waves to the subject, and a structure forming a layer in the subject. Means for measuring the thickness, means for measuring the speed of sound of the portion of the layer structure, and the layer thickness measured by the layer thickness measuring means and the sound speed in the layer structure measured by the sound speed measuring means. A delay time to be given to each transducer in consideration of an ultrasonic refraction effect in an ultrasound propagation path between the transducer and a designated focal point, and a refraction correction delay control to be supplied to the delay control means. Means, and a display unit for displaying an ultrasonic image. The sound velocity measuring means calculates the delay time error of each receiving channel by using the output of a delay circuit that performs phasing processing of the echo signals received by the plurality of transducers, and calculates the delay time error in the subject from the delay error. Includes sound speed measurement means for determining sound speed. Further, the sound velocity measuring means includes means for specifying a sound velocity measuring region to a fat layer of the subject and a tissue portion at a part of the fat layer, and measuring a sound velocity for each of the tissue layers.

 Further, the ultrasonic imaging apparatus of the present invention uses the parameters related to the ultrasonic probe including the lens layer thickness of the ultrasonic probe, the sound velocity of the lens layer, and the transducer array pitch, and the ultrasonic wave exits the transducer. A program to make a computer execute a method to calculate the delay time to be applied to each transducer taking into account the effect of refraction from when the lens reaches the specified focal point, or the lens layer thickness, the sound velocity of the lens layer Using the parameters related to the ultrasonic probe including the pitch between the transducers, and the data of the fat thickness of the subject, the sound velocity of the fat layer, and the sound velocity of the living tissue excluding the fat layer, the ultrasonic The program has a built-in program that allows a computer to execute a method of calculating the delay time given to each transducer taking into account the effect of refraction from exiting the transducer until reaching the designated focal point. are doing. BRIEF DESCRIPTION OF THE FIGURES

FIG. 1 is a block diagram showing a schematic configuration example of an ultrasonic imaging apparatus according to an embodiment of the present invention, FIG. 2 is a block diagram showing an embodiment of an ultrasonic beamformer without the ultrasonic imaging apparatus shown in FIG. 1, FIG. 3 is a diagram showing the relationship between the transducer array and the ultrasonic image, FIG. 4 is a diagram showing the relationship between the transducer array and the delay time during the transmission and reception of the ultrasound, and FIG. FIG. 6 is a flowchart showing a flow of tomographic image imaging to which an example of refraction correction is applied. FIG. 6 is a diagram showing an example of a method for measuring a layer thickness based on an image displayed on an image display unit. FIG. 7 is a diagram showing refraction of an ultrasonic pulse. FIG. 8 is a flowchart illustrating a first example of a procedure for calculating a delay time for refraction correction according to the present invention. FIG. 9 is a flowchart illustrating a procedure for calculating a delay time for refraction correction according to the present invention. FIG. 10 is a flowchart for explaining a second example, FIG. 10 is a diagram for explaining a relationship between a focal position and a transmitting aperture position in a phased array probe, FIG. 11 is an explanatory diagram of a refraction state of an ultrasonic pulse, and FIG. Illustration of the discrete Newton method, FIG. 13 is a flowchart illustrating a third example of the procedure of delay time calculation for refraction correction according to the present invention, and FIG. 14 is a procedure of delay time calculation for refraction correction according to the present invention. Flowchart illustrating a fourth example of FIG. 15 is an explanatory diagram of a refraction state of an ultrasonic pulse in a convex type probe, FIG. 16 is a block diagram showing a schematic configuration of another embodiment of the ultrasonic imaging apparatus according to the present invention, and FIG. FIG. 18 is a block diagram showing a schematic configuration of still another embodiment of the ultrasonic imaging apparatus, FIG. 18 is a graph showing a time required to reach a focal point from each transducer, and FIG. 19 is provided with a switch for a refraction correction function. FIG. 20 is a block diagram showing a device configuration for obtaining a sound speed for refraction correction, and FIG. 21 is a diagram showing a screen display example of sound speed measurement. BEST MODE FOR CARRYING OUT THE INVENTION

 Hereinafter, embodiments of the present invention will be described with reference to the drawings.

FIG. 1 is a schematic configuration diagram illustrating an example of an ultrasonic imaging apparatus according to the present invention. The ultrasonic imaging apparatus 100 includes a main body 10, an ultrasonic probe 20, and a cap. The main body 10 of the ultrasonic imaging apparatus includes a control unit 11, a memory 12, an image display unit 13, and an ultrasonic beam former 14. The control unit 11 is connected to an input unit 15 such as a keyboard and a pointing device. The ultrasonic probe 20 is detachable from the main body 10 of the ultrasonic imaging apparatus, and an appropriate probe is selected and attached according to a diagnosis target of the subject 30. The ultrasonic probe 20 includes a transducer array 21 and a memory 22. The memory 12 of the main body 10 stores a table of data unique to the probe corresponding to the ID number of the probe, and data such as a sound velocity in a subcutaneous fat layer and a sound velocity in a living tissue. The transducer-specific data referred to here includes, for example, the thickness of the lens layer, the sound velocity in the lens layer, and the transducer pitch. In the case of a hive or convex type probe, it refers to the radius of the probe, and in the case of an oblique type probe, it refers to parameters such as the oblique angle.

 The ultrasonic beam former 14 controls the operation of the transducer array 21 in the ultrasonic probe 20 under the control of the control unit 11. Ultrasonic image data is obtained by scanning an ultrasonic beam in a subject using this ultrasonic beamformer. The control unit 11 displays an ultrasonic tomographic image of the subject 30 on the image display unit 13 based on the obtained ultrasonic image data.

 FIG. 2 is a block diagram illustrating an example of an ultrasonic beamformer. The ultrasonic beamformer includes an operation control circuit 41, a pulser 42, a preamplifier and an AD converter 43, a delay circuit 44, and an addition circuit 45. In the ultrasonic beam former 14, a delay time is calculated and controlled by a calculation and control circuit 41. The arithmetic and control circuit 41 may be realized by a computer and software. In this case, a program for delay time calculation is recorded on a recording medium such as a ROM, and the calculation and control circuit 41 is realized by a configuration in which the program is read by a computer.

The beamforming mechanism for transmitting and receiving is as follows. First, each of the transducers 25 of the transducer array 21 included in the ultrasonic probe 20 from the arithmetic and control circuit 41. , 25 ,, 25 _2> ..., 25 _n , a pulse signal is sent at a timing shifted by the delay time of each oscillator. Each pulser 42 that receives the pulse signal immediately sends the pulse signal to the transducer 25 connected to itself. , 25 ぃ 25 ₂ ,. Send to Transducer 25. , 25 ₁ ; 25 ₂ ,..., 25 _n generate ultrasonic waves corresponding to the voltage due to their piezoelectricity, and transmit the ultrasonic waves into the subject 30. The ultrasonic waves reflected in the subject 30 return to the transducer array 21 of the ultrasonic probe 20 again, and the individual transducers 25. , 25 "25 ₂₎ ···, is converted into an electric signal by a piezoelectric at 25 _n, are amplified by preamplifier and AZD converter 43 is converted into a digital signal. The digital signal, the arithmetic and control circuit 41 The delay time is adjusted by the delay circuit 44 adjusted by the signal from the controller and the signal is added to the adder circuit 45 to be added, whereby the phasing addition is performed.

Next, a method of scanning the ultrasonic beam by the transducer array 21 provided in the ultrasonic probe 20 will be described. When scanning the ultrasound beam, as can be seen from the example of the ultrasound diagnostic image shown in Fig. 3, the ultrasound beam is scanned both in the depth direction and in the lateral direction. It is necessary to move the focal point of the sound beam. Here, in FIG. 3, if I is an ultrasonic image, the ultrasonic image I is a point F _nl , F _n2 , F _n3 , (where n is an integer from 1 to N) It consists of N received beam data. This point F _nl , F _n2 , F _n3 ,..., Is measured by a single ultrasonic beam. FIG. 4 is a diagram schematically showing an ultrasonic probe that transmits an ultrasonic pulse toward the focal point F „and detects a reflected signal from the focal point F„. Figure 4 shows how the wavefront moves at regular time intervals Ts. As shown in Fig. 4 (a), in order to focus the ultrasonic probe to F „, the ultrasonic pulse transmitted from the transducer 25 within the bore of the ultrasonic probe is used. An ultrasonic wave is transmitted with a delay time for each vibrator so that it converges at the focal point F „. In the case of Fig. 4 (a), transducers 25 located at both ends of the aperture. , _25π is 0, and the delay time Τ ”given to the j-th oscillator 25” is the oscillator 25 as shown. Is the time until a virtual ultrasound wavefront WA at position 25 _n reaches the oscillation surface of the vibrator 25 ". When an ultrasonic wave is transmitted with a delay time given to the operation of each transducer in this manner, the wavefront WA converges toward the focal point F „as shown in Fig. 4 (a). The wavefront WB of the reflected wave propagates as shown in Fig. 4 (b), and the oscillators 25 at both ends. , _25n, and arrives at the transducer surface of the transducer 25 ”with a delay time 時間”. Thus, in the child adds the phased signals by shifting the minute delay time in each of the ultrasonic probe diameter in the vibrator 25 ₀ to 25 _n, it is possible to pick up the reflected signal from the focal point F " .

 An image is acquired by performing this transmission / reception operation on point F (n is an integer from 1 to N). In an actual device, this transmission / reception operation is a method known as a dynamic focus method, that is, transmission is performed by focusing on a specific depth located in a region of interest (R0I: Region of Interest) within a subject. Transmits a sound wave, and changes the focal position of the received wave from shallow to deep, from region to region as the transmitted ultrasonic pulse travels through the subject during reception. Is used to obtain the received beam signal of When the measurement of one beam is completed, the beam position is sequentially shifted to the adjacent direction (azimuth direction), and the same transmission / reception operation is repeated to obtain N reception beam signals to form an image.

There are two ways to move the beam in the azimuth direction: one with a fixed aperture and the other with a variable aperture. There are two ways. The former is a scanning method used for sector-type probes. In this method, all transducers of the probe are used for transmission and reception every time, and the beam direction is changed radially. In this case, when the beam direction is in relation to the transmission aperture, in other words, the focal point of transmission and reception is slightly in front of the aperture. This has great significance in the calculation of the delay time later. The latter is a method used for a linear probe and a convex probe. This method differs from the former in that it does not use all transducers provided in the probe for one transmission.For example, a probe with a total number of transducers of 192 channels and a transmission of 64 channels in diameter Is used. In this case, when obtaining a beam at a certain position, transmission and reception are performed with the aperture positioned in front of the beam, and movement of the beam position is performed by moving the aperture, so that the focus is always in front of the aperture, contrary to the former case. How to transmit waves.

 Next, an example of the flow of tomographic image imaging using the ultrasonic imaging apparatus of the present invention will be described with reference to the flowchart of FIG. In the ultrasonic imaging apparatus, since the ultrasonic probe can be replaced by attaching and detaching itself, prior to imaging, in step 11, recognition of the ultrasonic probe 20 attached to the main body 10 of the ultrasonic imaging apparatus is performed. I do. This recognition is performed by reading a parameter such as a probe ID number stored in the memory 22 of the ultrasonic probe 20.

 Next, in step 12, a tomographic image of the subject is captured by the delay control data obtained based on the average sound velocity of the living body initially set in the apparatus without performing the refractive index correction described later. I do. FIG. 6 is a schematic diagram of a diagnostic screen displayed on the image display unit 13 at this time, and an image of a somewhat blurred target is displayed because no correction is applied. The image shown in Fig. 6 is an image taken by a convex type probe, and a subcutaneous fat layer is shown on the side in contact with the probe.

In step 13, the diagnostician measures and inputs the thickness of this fat layer. Specifically, the measurement is performed using a caliper conventionally provided in an ultrasonic imaging apparatus. In Fig. 6, the cursor of the caliper is positioned at the point where the fat layer indicated by A starts and the point where the fat layer indicated by B ends. Is determined. (In the example shown, 13 corrupt) This value can be used as the thickness of the fat layer. Fatty layer thickness entered by the diagnostician After that, all processes are executed in the diagnostic device.

 Hereinafter, the imaging mode is entered, and the diagnostician instructs the beamformer 14 from the input unit 15 via the control unit 11 of the ultrasonic imaging apparatus to the beamformer 14 in step 14. The beamformer 14 calculates the delay time after the refractive index correction described later has been performed by the arithmetic and control circuit 41 in step 15. The delay time calculated here is the delay time given to each transducer for transmission corresponding to the focal depth of the transmission, and the delay given to each transducer to form one ultrasonic reception beam. This is the time, that is, the delay time for continuously changing the focal position of the received wave as described above.

 Subsequently, in step 16, when the transducer array of the ultrasonic probe is excited using the determined transmission delay time, the ultrasonic array is converged to the focal point specified on the first ultrasonic beam line. An ultrasonic beam is transmitted. Next, in step 17, the received signal from the transducer array of the ultrasonic probe is phased using the receiving delay time calculated in step 15, and added to obtain the first received beam signal. can get.

 In step 18, it is determined whether signal acquisition from all points in the imaging screen has been completed. That is, when an ultrasonic image is formed by N ultrasonic beams, it is determined whether an ultrasonic beam of the Nth address has been obtained. If not completed, the flow returns to step 14 to change the beam position and perform transmission / reception. This is repeated, and when the scanning of the entire imaging range has been completed for the scanning in the depth direction and the scanning in the horizontal direction, the diagnostic images have been captured. After taking the diagnostic image, proceed to step 19, display it on the image display section 13, and end. In the image display, instead of displaying all the data for all points of a single diagnostic image after completing the data collection, each time data for one received beam is obtained in step 17, it is displayed as an image. You may make it display in the part 13 sequentially. This is possible by utilizing the function of the digital 'scan' converter provided in the ultrasonic imaging device.

Next, the processing content of “calculate the delay time after performing the refractive index correction” in step 15 of the flowchart in FIG. 5 will be described in detail. Here, as an example, a case where the transmitting surface of the probe has no curvature in the array direction, such as a linear sector type probe, will be described. For other types of transducers such as convex type, It can be handled in the same way, only the form of the expression changes.

 Referring to FIG. 7, consider a case where an electric pulse signal is sent to a vibrator 25 "having an aperture and then an ultrasonic wave emitted from the vibrator 25" reaches a focal position. The ultrasonic probe is provided with a transducer row arranged one-dimensionally at a transducer pitch p, and a lens layer having a thickness arranged in front of the transducer row. The lens layer is for converging the ultrasonic waves generated from the transducer row in a direction perpendicular to the transducer row (a direction perpendicular to the plane of FIG. 7). It is assumed that the sound velocity in the living tissue existing inside the fat layer of the living body is uniform on average, and that the fat layer has a certain thickness within the transmission aperture of the ultrasonic probe. Think. In this case, the ultrasonic wave transmitted from the transducer propagates through the three layers of the lens layer, the fat layer, and the living tissue to reach the focal point, and the refraction becomes a three-layer problem.

Now consider the path from the j-th transducer 25 '' located at a distance X from the center of the transducer row to the focal point located at depth F in front of the transducer row center. . The electric pulse signal is converted into ultrasonic waves by a piezoelectric vibrator 25 "having a matching layer on the front surface. The ultrasonic waves generated by the vibrator 25 '' travel through the lens layer of thickness by a distance _{Xl in} the array direction of the vibrator row, enter the subcutaneous fat layer at an incident angle, and at the boundary between the lens layer and the fat layer. The refracted (refraction angle of 0 ₂ ) ultrasonic wave travels through the thick subcutaneous fat layer by a distance x _{2 in the} direction of the array of transducer rows, and enters the living tissue below the fat layer at an incident angle θ ₂ I do. The ultrasonic wave refracted at the interface between the fat layer and the living tissue (refraction angle 0 ₃ ) travels in the living tissue in a diagonal direction of 距離_{3 in} the array direction of the vibrator row and d _{3 in} the depth direction. Reach the focus within.

The refraction of ultrasonic waves at the interface between each layer follows Snell's law. Therefore, assuming that the sound speed in the lens layer is _Cl , the sound speed in the fat layer is c ₂ , and the sound speed in the living tissue is c ₃ , Equation 1 is established between the refraction angle 0 ぃ θ, and the above.

C,

 (1)

sin Θ _i sin θ ₂ sin θ 3 Here, the sound speed C, is determined by the material of the lens, and can be measured. It is known. In the following description, the sound speeds c ₂ and c ₃ are also empirically assumed to be known as approximate values, and the following description is advanced. An embodiment will be described later. Incidentally, speed of sound _Cl in the memory 22 of the ultrasonic probe 20, also, the sound velocity c _2, c ₃ are assumed to be held in memory 12 of the ultrasonic imaging apparatus main body 10.

 In the present invention, two methods are proposed as numerical solutions to this problem. First, the first method will be described below. Erasing sin from Equation 1 gives Equation 2 below.

^ "X, ² ten _{^{= 2 + ά = ^ x 3}} 2 dozens ² ... (2) to force, Equation 2, can not be solved analytically in this state. In this case, seek an analytically solution Therefore, we solve Equation 2 with three simultaneous equations: The variables left behind when the variables are dropped are the lateral displacement _Xl in the lens layer and the lateral displacement x in the fat layer x _2, there is a choice about the or leave any lateral movement amount x ₃ in a living body, a most advantageous and Do Runowa chi ₃ under a number of conditions when considering the magnitude of the error This is because the speed of sound gradually increases in the lens layer, fat layer, and living tissue, so that the path of the sound wave passing through each layer gradually lays down, and under many imaging conditions. Therefore, in addition to this, the thickness is the largest, so in the actual system, the sound path in the living tissue has the largest lateral movement. Kikunaru -.. Since turn seek directly computing a large amount of Ru method der to reduce most the relative error, in the normal imaging conditions become better by obtaining the x ₃ and teeth force, regardless of the condition the lens layer thin top, thickness imaging conditions thickness and body tissue of the fat layer in relation to the thickness of the focal length of different is the depth of the imaging site. Yotsute fat layer, the x ₂ There may be an algorithm that considers whether to leave or leave x ₃ and then solve a polynomial of one variable, but here, the explanation of the method is omitted and only the method of obtaining x ₃ will be described. I do.

Eliminate X x ₂ from the simultaneous equations to obtain Equation 3. In Equation 3, we find X Therefore, in the following, this is simply abbreviated as x, and replaced with r ₂ = (c ₃ X c ₃ ) / (c ₂ X c ₂ ).

H (x) = ((r ₂ -l) x ² + r ₂ d ₃ ² )) c ₃ (X x) J x * ten d ^ —c ^ x (X .x) ²⁰ ten = 0

 H (x) H (x)

 ... (3) The calculation unit of the ultrasonic beamformer in the actual ultrasonic imaging device uses the MPU or DSP as described above. When DSP is used, all calculations are converted to sum of products. In a typical DSP, the square root is equivalent to 32 sums of products. Since the division depends on the value, the number of calculations is not known before the calculation. Therefore, transforming Equation 3 into a form without square root and division is important when using a DSP, and Equation 4 is the result of the operation.

I (x) = ((c , 2 one c ₃ ²⁾ x_ one c ₃ ² d ₃ ²⁾

(I (x) H (x) (X-x) ² + H (x) x ² d ₁ ² c ₁ ² + I (x) x ² d ₂ ² )… (4) 1 4 (X—x) ² x ² d ₂ ² I (x) ² H (x) = 0 According to Galois theory, it is known that no solution formula exists for polynomials whose order cannot be reduced by factorization if the degree is 5 or higher. ing. This makes it impossible to solve Equation 4 immediately. However, although Equation 4 is a 12th-order equation, it is a rational polynomial and is differentiable in the entire range, so it can be approximated as a straight line sufficiently near the solution. It is only necessary that X satisfy the condition that it is always near the solution.

And try to analyze the behavior of the formula 4 in the vicinity of the solution, but not to both sides square to remove the square root and have your conversion from Equation 3 to Equation 4, which _gl the preceding formula 3, after When the section and _{g 2, gl (x) -} g 2 (x) is g ₁ of solution were looking for a _{(x) X g 1 (x} ) -g 2 (x) X g ₂ (x) = (g, (x) — g ₂ (x)) x (g ₁ (x) + g ₂ (x)) , the original _gl (x) - in addition to the solutions of g ₂ (x), 8, also includes solutions of (χ) + (χ). This two-sided square operation is performed twice, which means that there are two extra solutions that are not relevant, but the solution newly included in Equation 4 by the second square is sufficiently original Since this is far from the solution, it does not matter here. Equation 4 has an inflection point near the solution because it also has a solution of _gl (x) + g ₂ (x), so the position where the approximate value is found needs to be closer to the true solution than this inflection point. It is. Using x + pXd ₃ ZF as the starting point of the approximate solution, using the notation in the figure as appropriate for that condition. An X at the vibrator position N and we do x _N is determined by the following equation 5. From now on, the left side of Equation 4 is expressed as f (x). d ₃ f (x _N — _t ten pd ₃ / F)

UPF f (x _N , + pd, / F), ⁵⁾ A method of calculating the delay time of each transducer in real time using this recurrence formula will be described with reference to a flowchart. First, using the flow chart in Fig. 8, a probe such as a linear probe or a convex probe, whose transmission plane is orthogonal to the line segment connecting the focal point and the center transducer of the aperture, is used. The case of calculating the delay time will be described. Under this condition, since the focal point is in front of the aperture, the delay time of each transducer is symmetrical within the aperture. In other words, it is enough to calculate only one half.

The parameters for the probe have already been given in step 11 of FIG. 5, and the value of the depth of focus has already been given in step 14 of FIG. Here, the delay time is calculated one by one toward the outside, starting from the transducer at the center of the aperture. First, in step 21, the object to be calculated is the central oscillator. Ultrasonic waves from the center transducer toward the focal point are perpendicularly incident on the layer interface and do not refract at the interface. = 0. Also, let the oscillator number N be 1 (the next oscillator with respect to the center oscillator). Then, in step 22, x. = 0, X! For oscillator number N = l! Is calculated by Equation 5 as follows. Next, proceeding to step 23, the delay time τ for the vibrator with the vibrator number N = l is calculated by the following series of formulas, that is, formula 6, based on the value obtained in step 22. e ₃ = arctan (x Zd ₃ )

Θ ₂ = arcsin (c ₂ sin θ ₃ , C ₃ )

θ ^ arcsinC sin θ ₃ main _{C 3) (6) x 2} = d 2 tan ^ 2

Α .-- X-- X ₂

Te - x tens phi ² + ₂ ² tens dc ^₂ + x _{3 2} tens d ₃ _z c 3 then proceeds to step 24, the transducer number is incremented by one, moved to the vibrator the adjacent, step Step 25 Returning to 22, x ₂ is calculated by the recurrence recurrence formula (Equation 5), and the delay time τ with respect to the oscillator of the oscillator number 2 is calculated based on the x ₂ by using Formula 6. This operation is performed for all the transducers having a diameter. If the determination in step 25 is "YES", it means that the delay time for all transducers has been calculated for this depth of focus, so the process proceeds to step 26, where the calculation result is output from the control circuit to the delay circuit. You. If the next depth of focus to be calculated is given to the ultrasonic beamformer, the same procedure is repeated for the all-diameter transducer. By repeating this for all the depths of focus, imaging is performed in which a shift due to refraction has been corrected.

Next, referring to the flowchart in Fig. 9, the line connecting the focal point and the center transducer of the aperture, such as the oblique type probe, the sector type probe, and the phased array type probe, is not orthogonal to the transmission plane. How to calculate the delay time for a type of probe, or the delay time for a type of probe in which the angle between the line connecting the focal point and the center transducer of the aperture and the transmitting surface changes The method for calculating the following is described. At this time, the relationship between the focal position, the aperture, and the vibrator row depends on the magnitude of the oblique angle 0, as schematically shown in FIG. (1) When the perpendicular foot A dropped from the focal point F to the plane where the transducers are aligned enters the transmission aperture

 (2) When A does not fall within the transmission aperture

There are two ways. In the latter case, consider a virtual vibrator row by cutting the distance from A to the vibrator at the end of the aperture at the element pitch. Based on this idea, before entering the calculation of the delay time, that is, referring to Fig. 5, in the next stage of step 13, the virtual oscillator at the position of A is calculated as the first term of the recurrence formula by Eq. Calculate X one by one to find X at the transducer at the end of the aperture. Parameters related to the probe, such as the oblique angle, have already been given in step 11 of FIG. 5, and values such as the depth of focus and the focus direction have already been given in step 14 of FIG.

 In step 31 of FIG. 9, for each depth and direction of focus within the area to be imaged, it is calculated whether the perpendicular foot (point A in FIG. 10) dropped from the focal point to the transmission plane is within the aperture. I do. If it is within the caliber, proceed to step 39, calculate the number of transducers from the perpendicular foot to the right and left ends of the caliber, and in step 40, determine the transducer closest to the foot to N = l, x. = 0. Subsequently, in step 33, X is obtained by the recurrence formula (Equation 5), and in step 34, the process of obtaining the delay time τ from X by Eq. 6 is performed up to the end of the aperture while increasing the oscillator number (S26). , S27). In this case, since symmetry cannot be used, it is necessary to calculate from the oscillator of の = 0 to the end of the aperture on both sides as in Step 37, Step 41, and Step 40. This is because Fcos 0 is not always an integral multiple of the oscillator pitch.

If the vertical foot dropped from the focal point to the transmitting surface is not within the aperture, proceed from step 31 to step 32 to read X from the transducer at the end of the aperture. Since this point Α the starting point for calculations during the time delay, the focal length F is F cos 0, d ₃ by using the original F, d ₃ = F cos 0- one original needs to be replaced.

 Then x at the end of the caliber x. And N = l. Then step

In step 33, X is obtained by the graduation formula (formula 5), and in step 34, the delay time τ is obtained from X by formula 6. Next, the process moves to the next vibrator in step 35, and the processes in steps 33 to 35 are repeated until it is determined in step 36 that the vibrator has reached the end of the aperture. With this, the depth of focus and the focus direction are delayed by all transducers. Since the delay time has been calculated, the calculation result is output from the operation / control circuit to the delay circuit in step 38. If the next depth of focus and the direction of focus to be calculated are given to the ultrasonic beamformer, the same calculation is performed again for all the transducers within the aperture. By repeating this for all the depths of focus and the focal directions, imaging is performed in which deviation due to refraction is corrected.

By using the recurrence formula of the above formula 5, the approximate value of the delay time can be obtained with sufficient accuracy (1/10 of the center frequency). This method has a sufficient accuracy compared with the conventional approximation formula, and has no loop in the algorithm as compared with the method of obtaining the accuracy by using the iteration, so that the calculation speed is considerably faster. Together with this algorithm and the recent increase in the speed of the arithmetic processing unit, a real-time high-speed delay time calculation algorithm including refraction by a fat layer as in the present invention has been realized. Furthermore, it is significantly advantageous in that errors do not accumulate as compared with the calculation of a general grading formula. This is because even if X for finding an approximate solution contains an error every time, f (x) is a function that does not include approximation, so that X and f (x), f '(X) This is because the error at x _N _, is not reflected in x _N in Equation 5 because there is no error between them.

This idea is a DSP-specific solution, but if there is no constraint, the calculation becomes easier. Since Equation 3 can be regarded as a straight line near the solution, there is no need to calculate the slope from the derivative, and we use x + p X d ₃ / F, x + p as the two points that clearly sandwich the solution. , と し て The solution is obtained as the following equation as a dividing point. Therefore, the processing speed can be improved by using the recurrence equation (Equation 7) instead of the recurrence equation (Equation 5). fix ^ + p d./F)

^Xn — ^{Xn i P} F tens ^{P (1} F) f (x _N , + pd ₃ / F) + f (x _N , + p)

(7) Next, another embodiment of the ultrasonic beam former according to the present invention will be described with reference to FIG. I will tell. In FIG. 11, it is assumed that the sound emitted from the transducer N at the transducer position X _N at an angle of 0 _N into the lens layer reaches the focal point F. When this angle 0 _N is obtained, the refraction angle in each layer is obtained, so that the sound path is determined and the delay time can be determined. Conventionally, to determine this delay time, sound is first emitted at an appropriate angle, and the side of the focal point through which the sound passes is calculated. If the sound is increased 0 _N if impassable farther than the focal, to reduce the 0 _N if impassable the near focus. By repeating this process until the sound passes through the focal point, 0 _N is obtained. However, in this method, the loop has to be repeated a considerable number of times before converging, and the calculation cannot be performed in the required speed. In order to achieve fast convergence, it is necessary to increase the manner in which θ _N is changed in accordance with the magnitude of the deviation. .

 The algorithm used in the ultrasonic beamformer according to the present embodiment is an improvement of the two-Juton-Raphson method, which can be applied to a system having no explicit function. That is, think as follows. First, assuming that an ultrasonic wave is emitted at an appropriate angle 0, the magnitude of the shift is ΔΧ (0). Since this Δ に な る becomes 0, 0 is the solution to be found, and in order to find the derivative of Δ ず れ, the deviation ΔΧ (Θ + d0) when emitted at an angle of Θ + d0 is found. As shown in FIG. 12, when a straight line passing through these two points crosses the X-axis in a space of 0 on the horizontal axis and ΔΧ on the vertical axis, Θ is a better approximate solution than the original 0. This can be expressed by Equation 8 as follows. —- _U J β h, 8 current)… f.ヽ new 'ΑΧ (θ ^-ΑΧ (θ ^ + άθ) As the shift amount, besides using ΔΧ, a method using ΔΥ or a method using the sum of squares of ΔΧ and ΔΥ can be considered. However, in this embodiment, a case where ΔX is used will be described.

By repeating this operation, 0 approaches the true solution. At this time, the smaller the difference between the first used 0 and the true solution, the less the repetition of the calculation, which is advantageous in terms of calculation time. In the present invention, as the initial value of 0, when calculating the delay time of the Ν-th oscillator, 0 _{N —} i obtained for the N−1-th oscillator is used. this In the method, it turned out that in practical calculation accuracy, it converged in two loops. Therefore, the number of loops is determined to be two. The number of loops is much smaller than the conventional method, so it is much better than the conventional method. When Θ is found for the Nth oscillator, this 0 is used as the initial value to find 0 for the N + 1st oscillator. In the case of this method as well, when the focal point is in front of the vibrator, there is no refraction, so that Θ = 0, which can be used as the starting point of the calculation.

 Using the flowchart in Figure 13, calculate the delay time of a probe, such as a linear probe or a convex probe, whose transmission plane is orthogonal to the line connecting the focal point and the center transducer of the aperture. A method for performing the above will be described. Under this condition, since the focal point is in front of the aperture, the delay time of each transducer is symmetrical within the aperture. In other words, it is sufficient to calculate only one half. The parameters for the probe have already been given in step 11 of FIG. 5, and the value of the depth of focus has already been given in step 14 of FIG.

 In step 51 in Fig. 13, the ultrasonic waves emitted from the transducer at the center of the aperture are perpendicularly incident on the interface between the lens layer and the fat layer and the interface between the fat layer and living tissue, and are focused without being refracted. Therefore, 0 = 0 and N = 1 for the vibrator with vibrator number 0. Next, proceeding to step 52, the deviation Δ Χ (Θ) between the path when ultrasonic waves are emitted from the transducer of the number N = l at an angle of 0 and the target focal position is obtained. Here, as the angle 0, the angle に 対 す る with respect to the next vibrator (vibrator of number 0) is used. In step 53, a deviation Δ Δ (0 + d0) between the path of the ultrasonic wave and the target focal position when the ultrasonic wave is radiated from the transducer of number 1 at an angle of 0 + d0 is determined.

Here, in the case of the convex type probe, unlike the linear type probe, it is calculated as follows. When observing an object with a convex probe, the probe is pressed against the object and observed, so the fat layer also bends along the probe. Assuming that the fat layer is concentric with the structure of the convex probe, the sound path is as shown in Figure 15. The two methods for calculating the deviation Δ when leaving the element and passing near the focal point and the method for finding the deviation Δ when leaving the focal point and passing near the element are basically the same. This system is not symmetrical because the reflection conditions and the sound path cannot go to the left half of Fig. 15, and the former method has good solution stability. Since it is not, it is desirable to solve by the latter method. If this sound path is analyzed for each layer and solved analytically, the position and angle of the point at which the sound enters the layer and exits the layer from the angle are calculated, and the end is reached by calculating the order of each layer The point is found, and the deviation Δ can also be found. For example, if the path of the sound in the living body is found outside the fat layer, the position of the sound's starting point is F = (X or y ₄ ), and the position at which the sound exits is P ₃ = (x ₃ , y ₃ ). Let R be the curvature of the layer interface on the outgoing side, and Θ be the angle at which sound enters. Then, x ₄ and y ₄ can be obtained by solving the always positive solution out of the ± 2 possible solutions of the quadratic equation that eliminates x ₄ from the simultaneous equation of Equation 9.

X twelve R ²

 (9)

χ ₃ -χ ₄

 tan ^

At this time, 0 ₄ can be obtained by considering the triangle 0FP ₃ . That is, since 0 _p3 is obtained from (x ₃ , y ₃ ) just obtained, it is obtained from the equation of 0 ₄ = 0 + 0 _p3 . After 0 ₃ Motomari with law theta ₄ or al Snell, determined the path as it passes through a certain layer sequentially in the same manner. In this example, the three-layer problem has been described as an example, but the same method can be used when there are more layers. Even if the interface between the fat layer and the living body outside the fat layer cannot be expressed by the concentric circle of the probe, if it can be expressed in the form of f (x ₃ ), use Equation 10, which is a modification of Equation 9, and θ ₄ 0 It can be obtained by using + arctan (-f '(x ₃ )) instead.

F (x ₃

χ, -χ.… ( ¹⁰⁾

 Two tan

Y ₄ -Y, Next, in step 54, a new 0 is obtained by extrapolation from equation 8. Subsequently, the process returns from step 55 to step 52, and uses the new 0 obtained in step 54 to shift ΔΧ (0). Ask for. In step 53, a deviation ΔΧ (Θ + d0) between the path of the ultrasonic wave when the ultrasonic wave is emitted at the new angle Θ + d0 and the target focal position is obtained. Next, in step 54, a new 0 is obtained by extrapolation from equation (8).

 Next, proceeding to step 56, the delay time τ is calculated from the obtained Θ by the following equation 11. Next, the oscillator number is increased by one, and the processing from step 42 is repeated. If it is determined in step 58 that the calculation has been completed for all the transducers within the working diameter, the process proceeds to step 59 and outputs the calculation result.

C ₂

Θ ₂ = arcsin (^-sin θ)

 ά d,

Te Λ Ha ten - - ² C + ³

c ^ cos Θ C ₂ COS θ c ₃ cos θ 3 The flowchart shown in Fig. 13 is based on the calculation of the recurrence formula in the flowchart shown in Fig. 8 and the calculation of the deviation from the focal position and the extrapolation. This is equivalent to what is replaced by the calculation processing of a new 0 by the above and the judgment processing of the number of loops for performing the loop processing of this part. Also in the case of the second method, when the focal point is not in front of the transmission aperture, the transmission aperture is determined in advance to 0 at the end vibrator before imaging, and that value is stored. A plurality of zeros at the transducer at the end of the transmission aperture are obtained for different subcutaneous fat layer thicknesses. When this condition is reached during imaging, this 0 is read and used to solve the recurrence equation as the first term of the recurrence equation. The flow of this calculation is shown in the flowchart of FIG. The details are easily explained by the fact that the relationship between the calculation process in FIG. 8 and the calculation process in FIG. 9 and the relationship between the calculation process in FIG. 13 and the calculation process in FIG. 14 are exactly the same in the first method. This method of pre-calculating the parameters for the transducer at the end of the aperture is a measure to prevent the DSP or MPU mounted on the current intermediate-level ultrasonic imaging device from having a margin, and a high-speed arithmetic processing unit can be mounted. Focus on high-end machines when calculating in real time It is also possible to perform all calculations in real time with imaging, including calculations that are wasted from the center of the virtual aperture at the position of the foot of the perpendicular that has fallen to the aperture plane. In that case, the calculation before imaging becomes unnecessary. Of course, in this case, it is not necessary to calculate the same vibrator spacing as the actual vibrator even for the virtual vibrator. It is valid. The calculation before the imaging needs to be performed every time when the thickness of the fat layer changes, so the imaging stops every time. There are two ways to solve the problem, as follows.

 In the first method, all calculations for each fat layer thickness are performed in advance within a conceivable range, and all calculation results before imaging are stored in the memory 22 of the ultrasonic probe 20 as data. It is a way to keep it. Then, when the ultrasonic probe 20 is connected to the ultrasonic imaging apparatus main body 10, the contents are transferred to the memory 12 of the ultrasonic imaging apparatus main body 10. During imaging, a delay time is calculated by the ultrasonic beamformer 14 with respect to this parameter and the depth of focus given by the control unit 11, and the delay time is given to the transducer array 21, and transmitted to the subject 30. I do. The reception is also phased by this delay time, and the control unit 11 calculates a diagnostic image and outputs it to the image display unit 13.

 On the other hand, in the second method, data calculated in advance is stored in a medium such as a CD-ROM and attached to the ultrasonic probe. The method is to install data from the CD-ROM or the like into the memory 12 of the ultrasonic imaging apparatus body 10 before or when using the probe. The operation after installation is exactly the same as the first method.

 FIG. 16 is a schematic configuration diagram showing another example of the ultrasonic imaging apparatus according to the present invention. In FIG. 16, the same functional portions as in FIG. 1 are denoted by the same reference numerals as in FIG. 1, and redundant description will be omitted. In the description so far, the input of the thickness of the subcutaneous fat layer has been performed by the diagnostician. That is, as shown in FIG. 6, when the diagnostician specifies a start portion and an end portion of the fat layer on the diagnostic screen of the diagnostic apparatus, the distance is displayed on the screen. This value was entered by the diagnostician.

The ultrasonic imaging apparatus shown in FIG. 16 includes a fat layer thickness calculation unit 16 in the control unit 11 so that screen output related to the subcutaneous fat layer can be directly input to the ultrasonic beamformer 14. It is something that has been done. In the fat layer thickness calculation mode, as shown in FIG. 6, when the diagnostician designates a portion A where the subcutaneous fat layer starts and a portion B where the subcutaneous fat layer starts by using a pointing device or the like, the fat layer thickness calculating section 16 calculates a distance between them. The calculation result is sent to beamformer 14.

 Next, a method for adaptively improving the image quality during imaging by utilizing the algorithm equipped with the ultrasonic beamformer of the present invention will be described. The calculation algorithm described so far assumed that the fat layer was homogeneous and constant in thickness, and that the sound velocity in the fat layer did not differ depending on the individual or the target site. This assumption was buried in a certain error in the delay time calculation, and the effect was not understood. However, if high-precision imaging is enabled by the present invention, it is clear that removing this premise improves image quality. In order to remove the premise, it is necessary to be able to change the sound speed of the fat layer slightly while taking an image. FIG. 17 is a schematic configuration diagram showing another example of the ultrasonic imaging apparatus according to the present invention. In FIG. 17, the same functional portions as in FIG. 1 are denoted by the same reference numerals as in FIG. 1, and duplicate description will be omitted. This ultrasonic imaging apparatus makes it possible to fine-tune the parameters relating to the fat layer, and to cope with a fat layer having a uniform thickness and non-constant thickness. Specifically, the ultrasonic imaging apparatus having the configuration shown in FIG. 1 was provided with an input device and a parameter calculation unit 17 so that the sound speed of the fat layer held by the ultrasonic beamformer could be varied. In this apparatus, a parameter corresponding to an input value is calculated in an input device and a parameter calculation unit 17, and the value is input to the ultrasonic beam former 14.

 The effect of changing this parameter is explained as follows. First, considering the case where there is no refraction, the time required to reach the focal point from each transducer is calculated as in Equation 12. This is shown in the graph in Fig. 18. Focus from each transducer

Time to reach

D… (12) At this time, if one or more of the parameters in Expression 12, that is, one or more of the sound velocity, the pitch p, and the F are changed, the curvature of the graph changes as shown in FIG. This is the same in principle, even in the case where the refraction of living tissue other than the lens layer, fat layer, and fat layer and the three layers must be considered, as in this case, the situation is complicated. Since the delay time is determined by the time required to reach the focal point from each transducer, the “vertical axis: the time required to reach the focal point from each transducer” versus the “horizontal axis: the position of each transducer” The delay time is represented by a relationship curve. Therefore, as described above, when the sound velocity of the fat layer is not constant depending on the location, this curve is distorted, and by changing various parameters so as to make a close fitting to the curve, the defocus can be reduced. It can be improved. It is most preferable to adjust the curve fitting so that the imaging screen is optimized, and the operability is good if the parameters are controlled by an external knob. At this time, it is also effective to adjust the thickness of the fat layer or the pitch of the vibrator instead of the sound velocity of the fat layer by appropriately interlocking and changing these parameters. In Fig. 18, (a) shows the case where the oscillator pitch and the sound velocity of the fat layer are not changed, and (b) and (c) show the case where the oscillator pitch and the fat layer sound velocity are adjusted.

 Hereinafter, an embodiment will be described for an example in which a layer structure exists on the surface of the subject. First, a specific example of obtaining the thickness and sound velocity of the layer structure on the surface of the subject during a medical examination will be described.

FIG. 19 shows an embodiment in which the examiner can select them. The table having the thickness shown in Fig. 19 (b) and the table of sound speed are prepared in advance and installed in the device. The examiner observes the image 170 displayed on the monitor 120, and obtains the image by referring to the depth scale 160. As a result, the thickness of the layer is read from the force, and a similar thickness is selected with switch SW2 (140) or the like, and the sound speed can be similarly selected with switch SW1 (150). In this case, switch SW1 and switch SW2 can be operated independently, so that the examiner can select values that are appropriate for each value. Regarding the sound speed, a numerical value may be used, or a selection may be made between a fat layer and a muscle layer. For example, if the fat layer 170 can be read as approximately 2 cm from the depth scale 160 shown on the screen of FIG. 19, 2 cm is selected by the layer thickness setting switch 140 provided on the console 130. Then, the screen Thickness is displayed at the top as 2cm. Next, the sound speed is set by operating the sound speed setting switch 150 in a predetermined direction. It is advisable to display those input values on the screen to confirm the selected values. The indication may be a numerical value such as 1450m / s, but may be an expression such as hard muscle, muscle, normal, fat, or high fat. Also, the switch can be a rotary type or anything. Of course, a touch panel may be used.

 The above description has been given of the embodiment in which the influence of the refraction of the ultrasonic wave generated from the layer structure of the living body is removed and the image quality of the ultrasonic image is improved, but in order to further improve the image quality of the ultrasonic image, It is necessary to accurately grasp the propagation speed of ultrasonic waves in a living body and reflect the data in the above-mentioned refraction correction control by an ultrasonic beamformer.

 For this reason, FIG. 20 shows another embodiment of the present invention. In this method, the sound speed of a living body having a layer structure is obtained, and the refraction correction processing of the above-described embodiment is performed based on the sound speed. FIG. 20 is a block diagram showing an embodiment of a part for obtaining the sound speed in the ultrasonic imaging apparatus. In the figure, reference numeral 200 denotes a digitally controllable digital delay unit that delay-controls a plurality of ultrasonic signals received by the probe and outputs a reception beam signal, and has a circuit having a number of channels corresponding to the number of transducers used for reception. have. Reference numeral 210 denotes a delay data error estimator that inputs a plurality of signals delayed by the digital delay unit 200 and estimates an error of the digital delay data subjected to delay control with respect to true delay data by calculation, and 220 denotes a digital delay unit A digital delay control unit that controls the operation of each of the 200 channels.

In addition to the above configuration, the ultrasonic imaging apparatus of the present embodiment further includes a sound speed corresponding delay time recording unit 230 that stores delay times due to a plurality of medium sound speeds in advance, and a delay obtained by the delay error estimation unit 210. A new delay time is calculated from the error, and the calculated delay time is compared with a value stored in the sound speed corresponding delay time recording unit 230, and a delay time comparing unit 240 which outputs data closest to the stored value, and a sound speed corresponding delay The sound speed data recording unit 250 that records the medium sound speed based on the delay time data stored in the time recording unit 230, and the delay time from the delay time recording location that matches the output of the delay time comparison unit 240 A medium sound speed selection unit 260 for selecting the medium sound speed with reference to the recording unit; The output line of the medium sound speed selection unit 260 is connected to the arithmetic and control circuit 41 shown in FIG. The digital delay control unit 220 is connected so as to be controlled by the arithmetic and control circuit 41.

 As described above, in the configuration shown in FIG. 2, the delay circuit is digitally controllable, a delay error estimator 210 is provided at the output thereof, and a digital delay controller 220 is added to the arithmetic and control circuit 41. Further, as described above, by newly providing the sound speed corresponding delay time recording unit 230, the delay time comparing unit 240, the sound speed data recording unit 250, and the medium sound speed selecting unit 260, the ultrasonic imaging apparatus of the present embodiment can realize the true medium sound speed. Refraction correction with a sound speed almost equal to that of the above becomes possible.

 Next, the operation of the portion configured as described above will be described. First, delay time data obtained by assuming the sound velocity in the living body to be, for example, the equal sound velocity of the average value of the living body is calculated and output from the control circuit 42 to the digital delay unit 2 and the ultrasonic wave is transmitted into the subject. . At this time, the focal point of the transmission is set to an appropriate depth. Then, the delay time data D based on the same sound speed as the transmission wave is supplied from the calculation / control circuit 41 to the digital delay unit 200 via the digital delay unit 220, and the received signal is delay-controlled.

Then, a signal which is a delay-controlled signal of each channel of the digital delay unit 200 and which is not subjected to addition processing for phasing is input to the delay error estimation unit 210, and calculates an error with respect to the delay time data D used. , And outputs D _cl = D + AD. As the calculation method, the correlation processing method disclosed in the above-mentioned Japanese Patent Application Laid-Open No. 8-317923 can be used. D _cl output from delay error estimation section 210 is input to delay time comparison section 240. Then, the delay time comparing unit 240 compares the data recorded in the sound speed corresponding delay time recording unit 230 with the input data D _cl , selects the delay time data closest to D _cl, and sends it to the medium sound speed selecting unit 260. Output. When the data is input, the medium sound speed selection unit 260 selects the sound speed of the input data. This selection can be made by associating the storage address of the data of the sound speed corresponding delay time recording unit 230 with the sound speed data of the sound speed recording unit 250. Therefore, the sound speed corresponding delay time recording unit 230 and the sound speed recording unit 250 can be integrated into one. ,

The sound speed data selected by the medium sound speed selection unit 260 is fed back to the digital delay control unit via the arithmetic and control circuit 41. This feedback circuit Is for repetition of the above operation in order to obtain an accurate sound speed. If necessary, the arithmetic and control circuit 41 issues a command to repeatedly execute the above operation. With the above operation, the sound velocity in the living body can be measured as an estimated value.

Note that, in order to simplify the comparison operation in the delay time comparison unit 240, the data D _cl output by the delay error estimation unit 210 and the data recorded in the sound speed corresponding delay time recording unit 230 are two-dimensional distribution data. It is useful to be able to use the curve fitting method because the amount of information to be handled can be reduced. Furthermore, it is useful to obtain the difference of the delay time distribution and to fit a first-order straight line to the difference delay time sequence because the amount of information to be handled is further reduced.

 Next, a method of obtaining the sound speed of each layer of the living body having the layer structure with the configuration shown in FIG. 20 will be described. FIG. 21 is an ultrasonic tomographic image of the subject displayed on the monitor of the ultrasonic imaging apparatus. In the image, 170 is a fat layer, and 190 is a tissue region deeper than the fat layer 170. First, the sound speed of the fat layer 170 is determined. For this purpose, the focal position is set using the above-mentioned cursor of the caliber at an appropriate depth position inside the fat layer 170, for example, a measurement position near the boundary between the fat layer 170 and the tissue region, and the value is measured. A signal for giving the focal position is supplied to the arithmetic / control circuit 41. After that, an ultrasonic pulse is transmitted from the probe and the reflected signal is received. In order to determine the sound speed of the fat layer 170, it is necessary to take only the reflected signal from the fat layer into the sound speed measuring unit shown in FIG. For this purpose, the arithmetic control circuit 41 is provided with a gating function. The gating function may be linked to the caliber function. Then, the reflected signal from the fat layer 170 is taken into the digital delay unit 200 by this gating function, and the sound speed is obtained in accordance with the operation description of the configuration shown in FIG. The desired sound velocity may be at a certain point, but it is desirable to determine the sound velocity at multiple points and then determine the average sound velocity from them.

Next, the sound velocity in the tissue region deeper than the fat layer 170 is determined. At this time, the focal position for measurement is set to the position 190 shown in Fig. 21. The above-mentioned caliber function is also used for this setting. Then, a signal for giving a focal position is supplied to the arithmetic and control circuit 41, and thereafter, an ultrasonic pulse is transmitted from the probe and a reflected signal thereof is received. In order to find the sound velocity in the tissue area, it is necessary to capture only the reflected signal from the area However, this is not possible because the reflected wave in this case must pass through the fat layer 170. Therefore, a reflection signal from the focal position is captured by the above-mentioned gating function. Then, the speed of sound is obtained from the reflected signal. The sound speed obtained here indicates the average sound speed between the probe surface and the measurement point.

The sound velocity in the tissue region can be obtained from the above two measurements. The average speed of sound was measured first fat layer, then the obtained fat layer and the average speed of sound including both tissue region put c _a and code. Then, assuming that the thickness of the fat layer, that is, the depth of the focal point position 180 is l _f and the focal depth of the latter measurement is l _a , the sound velocity c ₂ in the tissue region can be obtained by Expression 13.

( 13)

c, thereby the speed of sound Motoma' fat layer and the acoustic velocity c ₂ of the thread! ^ middle by applied to the following equation 1 above, the lens layer of the ultrasonic wave, in consideration of the speed of sound fat layer and tissues refraction Correction is possible. The above measured values are displayed as numerical values on the monitor screen. In the display example in Figure 21, V: 1450 / 1540m / s indicates that the sound velocity of the fat layer is 1450m / s and the sound velocity in the tissue is 1540m / s, and t: 2 / 8cm indicates that the thickness of the fat layer is 2cm and the tissue It indicates that the focal position of is 8cm deep.

Next, another embodiment will be described. This embodiment considers the correction of the distance measurement function of the carrier used for the sound velocity measurement. That is, the distance measurement function of the caliper incorporated in the ultrasonic imaging apparatus is performed by an operation based on the speed of sound initially set in the apparatus. Therefore, it is desirable that the distance measured by the caliper of l _f , 1 _β used in Equation 13 be used after being corrected to the value based on the actual sound speed. The present embodiment corresponds to this. In this embodiment, first, the ultrasonic imaging apparatus is driven to acquire an ultrasonic tomographic image of a section including the region of interest (R0I), and the tomographic image is displayed on a monitor. In this state, the refraction correction execution switch 310 arranged on the operation panel of the ultrasonic imaging apparatus shown in FIG. 19 is turned on. This refraction correction execution switch 310 performs refraction correction when turned on, and refraction when turned off. This enables normal imaging without correction. When this switch operation is performed, a caliper 300 consisting of two points of force is displayed on the screen of the monitor 120. Then, by operating an input operating device such as a trackball or a joystick, one cursor of the caliper 300 is moved to the body surface of the subject, and the other cursor is moved to the end of the fat layer to fix the input information. Operating the key (Enter key) 320 specifies the fat layer to be measured. Next, of the two cursors in the caliper 300, the force sol positioned at the end of the fat layer is moved into the region of interest at a point deeper than that, and the sound velocity in the tissue is operated by operating the key 320. Identify measurement points for measurement. The data of the measurement points input by the above two-step operation are read into the control unit 11, and the calculation of the above-described sound velocity measurement method is executed.

The device setting sound speed is v. , Its sound velocity V. Caliber Depth at x. , Assuming the measured average sound velocity of the fat layer c _fn , the true thickness l _ft of the fat layer can be obtained from Eq.

1ft― ¾ "Cfm, V. (14)

These values _CfB , i _ft from the true average sound velocity _c in the tissue. _a , the true sound propagation distance in the tissue l. By calculating _t and applying it to Equation 13, delay time data for refraction correction can be calculated, and a good image can be obtained by transmitting and receiving ultrasonic waves using the focus data.

 Various methods other than those described above can be used to determine the sound speed and distance of a living body having a layered structure. As an example, focusing on a specific beam of the ultrasonic beam, for example, the center beam, detecting a place where the intensity of the received signal in the beam is very large near the body surface, obtaining the time as the time, The calculation may be performed automatically. This measurement method is possible if the threshold value for detection is determined because the reflected signal is large at the boundary of the layer structure.

If the refraction correction method described above is programmed and incorporated into the device as an automatic sequence, and repeatedly executed every few frames during imaging to update the set values, if there are multiple inspection sites Move the probe as shown The refraction correction is performed automatically, so the image is always good.

 In a device that enables refraction correction, a normal image is displayed on the left of the screen, and a right image is displayed on the left side of the screen using a method in which a displayed image is marked to distinguish between a normal image and a refraction corrected image. In which a refraction-corrected image is displayed. Even if the image at the point of interest is good, the image may be distorted in other regions in the case of a complex biological structure. Therefore, simultaneous display makes it possible to compare images with and without refraction correction in real time, which is effective for the user. It is also useful to specify the diagnosis site as R0I, acquire an image with refractive correction applied only to that site, fit it into a normal image, and display it.

 The present invention is not limited to the above specific embodiment, and various modifications can be made without departing from the scope of the technical idea.

 As described above, according to the present invention, the effect of refraction due to the lens layer and the fat layer (or the muscle layer, etc.) can be reduced, so that the image quality of an ultrasonic image can be improved. Further, according to the present invention, the sound velocity of each layer of the subject having a layered structure can be measured, and the influence of the refraction of the ultrasonic wave in each layer can be reduced by taking the value into account, thereby further improving the image quality. can do.

Claims

The scope of the claims

1. Ultrasonic probe with arrayed transducers and delay time control for each transducer to transmit or receive ultrasound when transmitting or receiving ultrasonic waves to / from the subject A delay control unit that performs the focusing of the transmitting or receiving wave by incorporating a refraction effect of the ultrasonic wave by the ultrasonic wave propagating medium between the arrayed transducers and the set focal position. An ultrasonic imaging apparatus comprising: a refraction correction delay data generating unit to be supplied to a control unit; and a display unit for displaying an ultrasonic image.

2. The delay control means previously stores delay time data obtained based on the average sound velocity of the living body, and is preceded by ultrasonic transmission / reception for obtaining refraction correction data using the stored delay time data. The ultrasonic imaging apparatus according to claim 1, wherein the ultrasonic imaging is performed.

3. The refraction correction delay data generating means uses parameters related to the ultrasonic probe including a lens layer thickness, a sound velocity of the lens layer, and an arrangement pitch of the transducer, and calculates a distance between the transducer and a designated focal point. 2. The ultrasonic imaging apparatus according to claim 1, wherein a delay time given to each transducer is obtained by calculation in consideration of an ultrasonic refraction effect in an ultrasonic propagation path.

4. The refraction correction delay data generating means includes: a parameter relating to the ultrasonic probe including a lens layer thickness, a lens layer sound speed, and a transducer pitch; a fat layer thickness and a sound speed of a fat layer of a subject; Calculating the delay time given to each transducer by taking into account the ultrasonic refraction effect in the ultrasonic propagation path between the transducer and a designated focal point, using the sound velocity data of 4. The ultrasonic imaging apparatus according to item 2 or 3.

5. The refraction correction delay data generation control means delays using a parameter gradually solved from a parameter relating to a sound path from the transducer next to the transducer to be calculated to the focal point. The ultrasonic imaging apparatus according to claim 1, wherein the time is obtained by calculation.

6. Means for detecting data related to the layer structure of the subject for refraction correction from the screen of the display unit on which the ultrasonic image is displayed, and using the output of the layer structure data detecting means to determine the layer structure of the subject based on the output. 2. An ultrasonic imaging apparatus according to claim 1, further comprising means for generating delay control data taking into account the effect of sound wave refraction.

7. The method according to claim 6, wherein the layer structure data detecting means displays two movable cursors on the screen, and includes a caliber for measuring a distance between the force and the sol on the screen. An ultrasonic imaging apparatus according to claim 1.

8. An ultrasonic probe with an arrayed transducer and a delay time for each transducer to perform transmit focusing or receive focus when transmitting or receiving ultrasonic waves to and from the subject. Delay control means for controlling; means for measuring the thickness of a layered structure in the subject; means for measuring the speed of sound of a portion of the layer structure; and a layer measured by the layer thickness measuring means. Using the thickness and the sound velocity in the layer structure measured by the sound velocity measuring means, each vibration is considered in consideration of the ultrasonic refraction effect in the ultrasonic wave propagation path between the vibrator and the designated focal point. An ultrasonic imaging apparatus, comprising: a refraction correction delay control unit that determines a delay time to be given to a child and supplies the delay correction unit to the delay control unit; and a display unit that displays an ultrasonic image.

9. The sound velocity measuring means calculates the delay time error of each reception channel using the output of a delay circuit that performs phasing processing of the echo signals received by the plurality of transducers, and calculates the delay time error in the subject based on the delay error. Sound velocity measuring means for determining the sound velocity of the object An ultrasonic imaging apparatus according to claim 8.

10. The sound velocity measurement means includes means for specifying a sound velocity measurement region to a fat layer of a subject and a tissue portion inside the fat layer, and measuring a sound velocity for each of the fat layers. 10. The ultrasonic imaging apparatus according to 9.

11. Using the parameters related to the ultrasonic probe including the lens layer thickness, the sound velocity of the lens layer, and the transducer array pitch, the ultrasonic wave is refracted from the transducer before reaching the designated focal point. An ultrasonic imaging apparatus comprising a program for causing a computer to execute a method of calculating a delay time given to each transducer in consideration of an effect.

12. Parameters of the ultrasound probe, including the lens layer thickness, the lens layer sound speed, the pitch between transducers, and the fat layer thickness of the subject, the sound speed of the fat layer, and the sound speed of living tissue excluding the fat layer. And using the data to calculate the delay time to be given to each transducer in consideration of the effect of refraction from the time the ultrasonic wave leaves the transducer to the point at which a designated focal point is reached. An ultrasonic imaging apparatus characterized by incorporating a program.