KR101306491B1 - A method for measuring shear modulus of tissue - Google Patents

Abstract

PURPOSE: A method for measuring the transverse elasticity characteristics of tissues is provided to estimate quantitatively the speed of transverse waves which occurs perpendicularly to the transmission direction of ultrasonic waves, thereby improving the accuracy of diagnosis results. CONSTITUTION: An ultrasonic image apparatus transmits ultrasonic waves to a first phantom changing a length of a transmission time multiple times. The apparatus measures the maximum displacement time when the displacement of the first phantom is maximized along a transmission direction in a focus depth of the ultrasonic waves. The apparatus obtains the transverse elasticity characteristics of the first phantom using the known specific parameters of the apparatus and correlation between the transmission time and the maximum displacement time. The transverse elasticity characteristics of the first phantom are the acoustic velocity of transverse waves which occur perpendicularly to the transmission direction of the ultrasonic waves.

Description

A Method for Measuring Shear Modulus of Tissue

The present invention relates to a method for measuring lateral elasticity characteristics from deformation generated in human tissue by acoustic radiation force when transmitting high power ultrasound in a medical ultrasound imaging apparatus, and more particularly, transverse elasticity from deformation caused by ultrasonic transmission. Deformation generated in the human tissue during the transmission of high power ultrasound by using the unique parameters of the ultrasound imaging apparatus obtained from a test phantom whose characteristics are known in advance and displacement characteristics of the human tissue to be diagnosed by the ultrasonic transmission. The present invention relates to a method for measuring transverse elastic properties of human tissues from which the transverse elastic properties can be quantitatively and easily measured.

In general, lesion tissues such as cancers or tumors that occur in the soft tissues of the human body have harder properties than those of surrounding tissues. Can be.

Medical elastic imaging, which is widely commercialized in recent years, is based on the mechanical properties of soft tissues of human body, and measures elasticity of tissues by using deformation of tissues generated when external force is applied to the diagnosis site. It is a technique of detecting lesion tissue by imaging.

In the medical elastic imaging method according to the prior art, the method using the compressive mudulus characteristic of the tissue was mainly used. This is performed by the diagnoser by compressing the diagnosis site by directly manipulating a transducer that generates ultrasound by hand. The external force was applied, and the strain of the tissue generated at this time was measured by imaging with ultrasound.

As described above, the elastic imaging method using the longitudinal elasticity characteristic of the tissue at the diagnosis site is disclosed in detail in the following Document 1, etc. filed by the inventors of the present invention.

However, the method using the elasticity of these tissues is useful for imaging areas close to the skin such as the breast or prostate, but it is not easy to provide external compressive force to areas located inside the human body such as the liver or kidneys. There was a problem that it is difficult to accurately diagnose the occurrence of tissue.

Therefore, recently, when transmitting a high-power ultrasound to the human tissue inside the human body to determine the occurrence of the lesion tissue by using the elastic characteristics that are deformed by being pushed in the direction of the ultrasonic sound field transmitted by the acoustic radiation force generated from the focus A technique using acoustic radiation force impulse imaging has been developed.

However, even when measuring the elastic properties of the tissue by using the acoustic radiation imaging method as described above only because the elastic properties of the tissue can be known only qualitatively, there is a problem that the accuracy of the diagnosis result is reduced.

Therefore, there is an urgent need for a technique capable of improving the accuracy of a diagnosis result by quantitatively measuring elastic properties of deformation of human tissue caused by the above-mentioned acoustic radiation.

[Document 1] Korean Patent Publication No. 2007-013981 (published Jan. 31, 2007)

The present invention is to solve the problems of the prior art as described above, an object of the present invention is a deformation that occurs in the sound field direction of the transmitted ultrasound inside the human tissue when transmitting a high-power ultrasound in the medical ultrasound imaging apparatus By measuring the transverse elastic properties of the human tissue by using a quantitative measure to provide a method of measuring the transverse elastic properties of the human tissue that can greatly improve the accuracy of the diagnosis results by the ultrasound imaging apparatus.

In order to achieve the above object, the method of measuring the lateral elasticity of human tissue according to the present invention transmits ultrasonic waves while changing the length of the transmission time τ a plurality of times to the first phantom to be diagnosed using an ultrasonic imaging apparatus. In each case, a first step of measuring a maximum displacement time (t _max ) at which the displacement of the first phantom is maximized in the transmission direction at the depth of focus of the ultrasonic wave, and the transmission time obtained from the measurement result of the first step and a second step of obtaining a lateral elasticity characteristic of the first phantom by using a correlation between the τ and the maximum displacement time t _max and an inherent parameter K of the ultrasonic imaging apparatus known in advance. It is done.

In addition, the second step may include a step 2-1 of approximating the maximum displacement time t _max from the measurement result of the first step to the first function of the transmission time τ, and the maximum value of the first function. And calculating a transverse elasticity characteristic of the first phantom using the intercept β of the displacement time t _max and the inherent parameter K of the ultrasound imaging apparatus.

In addition, the lateral elasticity characteristic of the first phantom is the sound velocity (c _s ) of the transverse wave generated in the transverse direction which is a direction perpendicular to the transmission direction of the ultrasonic wave by vibrating the first phantom by the transmitted ultrasonic waves, the second Step -2 is characterized by obtaining the sound velocity (c _s ) of the shear wave by the following Equation 1.

[Equation 1]

In addition, the step 2-2 further includes the step of obtaining a transverse elastic modulus (μ), which is a lateral elastic modulus of the first phantom, from the sound velocity (c _s ) of the shear wave according to Equation 2 below. It is characterized by.

[Equation 2]

In addition, the inherent parameter K of the ultrasonic imaging apparatus is transmitted to the second phantom in which the sound velocity c _s0 of the shear wave generated in the direction perpendicular to the transmission direction of the ultrasonic waves by vibrating by the ultrasonic transmission is known in advance. Ultrasonic waves are transmitted by varying the length of time τ _{0 a} plurality of times, and in each case, the maximum displacement time t _max0 at which the displacement of the second phantom is maximized in the transmission direction at the depth of focus of the ultrasonic waves is measured. After approximating the maximum displacement time (t _max0 ) from the measurement result to the first function of the transmission time (τ ₀ ), the intercept (β ₀ ) of the maximum displacement time (t _max0 ) with respect to the primary function is expressed by Equation 3 below. It is characterized by being obtained by application.

[Equation 3]

As described above, the method of measuring the lateral elasticity of human tissues according to the present invention utilizes the displacement characteristics of the human tissues according to the change in the length of time of the transmission sound field when generating an ultrasonic radiation force on the human tissue to be diagnosed by transmitting ultrasonic waves. By quantitatively estimating the velocity of the shear wave generated in the direction perpendicular to the transmission direction of the ultrasonic waves inside the human tissue, the transverse elastic properties of the human tissue can be measured quantitatively, thereby improving the accuracy of the diagnosis result by the ultrasound imaging apparatus. The advantage is that it can be improved.

1 is a view for explaining the principle of generating acoustic radiation and shear waves inside the human tissue by ultrasonic transmission;
2 to 7 are views for explaining the theoretical background of the method of measuring the transverse elasticity of the human tissue according to the present invention, respectively;
8 to 10 are views for explaining a method of measuring the transverse elastic properties of human tissue according to the present invention, and
FIG. 11 is a diagram illustrating the configuration of an ultrasound imaging apparatus to which a method of measuring transverse elasticity of human tissues is applied.

Hereinafter, preferred embodiments of the present invention will be described in detail with reference to the accompanying drawings.

1 is a view for explaining the principle of the generation of acoustic radiation and shear waves inside the human tissue by ultrasonic transmission, Figures 2 to 7 each show the theoretical background of the method of measuring the transverse elastic properties of the human tissue according to the present invention It is a figure for demonstrating.

8 to 10 are views for explaining a method of measuring the lateral elasticity of the human tissue according to the present invention, Figure 11 is a configuration of the ultrasonic imaging apparatus to which the method of measuring the lateral elasticity of the human tissue according to the present invention The figure shown.

In general, when a high-power ultrasound of about 1MHz to 10MHz is transmitted to the soft tissue of the human body, the tissue is pushed by the sound pressure of the ultrasound in the direction of transmission of the ultrasound. This force is called acoustic radiation force, and the magnitude thereof is as follows. It appears as shown in [Equation 1].

Where I is the sound pressure intensity [w / cm 2] of the ultrasonic wave, α is the attenuation coefficient of the tissue [Np / m], and c is the sound velocity of the longitudinal wave [m / s].

When the tissue is pushed out by the acoustic radiation force, as shown in Fig. 1, a transverse wave is generated in a transverse direction perpendicular to the transmission direction of the ultrasonic sound field due to the vibration of the tissue, as described above. When the lesion tissue is generated, the soft tissue becomes harder and the elasticity value increases, so the sound velocity of the shear wave increases in proportion to it.

In this case, the sound velocity (c _s ) of the shear wave is known to have a relationship as shown in [Equation 2] between the elastic modulus of the tissue, where E and μ are the elastic modulus of the tissue and Is the lateral elastic modulus, ρ is the density of the tissue and ν is the Poisson's ratio.

Therefore, when measuring the sound velocity of the shear wave generated by the acoustic radiation force in the human tissue to be diagnosed, the transverse elasticity which is the transverse elasticity value of the tissue by [Equation 2] (that is, μ = ρc _s ² ) By knowing the counts, it is possible to diagnose whether lesion tissue has developed in the tissue.

The acoustic radiation imaging method used in the present invention is a method for diagnosing the occurrence of lesion tissue by measuring and imaging the lateral elasticity characteristics of human tissues using the characteristics as described above. The apparatus means an imaging apparatus using the acoustic radiation imaging method by ultrasonic transmission.

The present invention uses an indirect method by estimating the transverse elastic properties (specifically, the sound velocity or transverse modulus of shear waves) of tissue using the displacement characteristics of the human tissue generated by acoustic radiation as described above. It is characterized by measuring quantitatively, hereinafter will be described in detail with respect to the theoretical basis of the measuring method of the transverse elastic properties of human tissue according to the present invention.

As described above, when transmitting ultrasonic waves of several hundreds of hours to generate acoustic radiation force, the tissue is pushed back to the original position. In this case, when the transmission time of the ultrasonic waves is increased, a relatively large acoustic radiation force is generated. The tissue is pushed further away and returned to its original position.

That is, the transmission time length τ of the ultrasonic sound field transmitted to generate acoustic radiation force in the elastic medium such as the soft tissue of the human body and the time t _max at which the medium reaches the maximum displacement have a linear proportional relationship. Looking at it as follows.

First, the general time waveform of the displacement generated by the acoustic radiation force at the focus of the focused ultrasound beam initially rises relatively quickly with time as shown in FIG. 2 and gradually decreases after reaching the maximum value. It shows a tendency.

At this time, the rise time of the displacement, that is, the time t _max to reach the maximum displacement is mainly determined by the elasticity of the tissue, and the relaxation curve in which the displacement decreases is greatly influenced by the viscosity.

In addition, the lateral elasticity value of the tissue is closely related to the velocity of the transverse wave generated in the direction perpendicular to the transmission direction of the ultrasonic wave by the vibration of the tissue.

In the present invention using this characteristic, the sound velocity of the shear wave can be quantitatively measured simply by measuring the t _max .

To this end, the present invention considers the spatial distribution of the volumetric force (focused ultrasound beam) as shown in FIG. 3 in order to analyze the displacement-time curve to determine the characteristics of the rise time. The volumetric force acts in the y direction and W _xz And W _y are the transmission sound field widths in the lateral and depth directions, respectively.

The displacement rise time in the elastic phantom is expressed by an approximation equation as shown in [Equation 3] below for the ellipsoidal shape force distribution whose pulse width is τ ₀ and the magnitude of the volume force is reduced to -20dB with the 3D Gaussian distribution. Can be.

Where the constants p and b are constants that depend on the spatial and shape distributions of the volumetric forces. The above equation is the displacement rise time at the origin, which is the center of the volume force distribution. Therefore, when t _max is measured experimentally, the shear wave velocity c _s can be calculated based on [Equation 3].

It is necessary to consider whether Equation 3 for the volumetric force distribution of the Gaussian spheroid form satisfies the other types of volumetric force. Since the shape is also different, it is a process for checking whether the shear wave velocity can be estimated using Equation 3 with an arbitrary ultrasonic imaging apparatus (that is, using the negative radiative imaging method).

The displacement-time curve according to the pulse width τ ₀ in the case of volumetric force of homogeneous hexahedron type with uniform and hexahedral distribution (see FIG. 3 (a)) with a volume of 1 N / m 3. Is shown in FIG.

As the pulse width increases, the maximum displacement u _max and the maximum displacement arrival time t _max increase. However, when the pulse width exceeds approximately 100 ms, the increase of u _max is slowed down, and when the pulse width is over 1000 Hz, the maximum displacement reaches saturation.

Referring to the dotted line of FIG. 4, in the region where the pulse width is small, t _max increases linearly with τ ₀ . The change in t _max according to the pulse width for the uniform hexahedron, homogeneous spheroid, and Gaussian ellipsoid is shown in FIG. 5.

The dotted lines in FIG. 5 are linear approximations of data satisfying linearity in the region of 2c _s τ ₀ <W _xz for given volumetric force distribution forms. Linear approximated constant values for the various volumetric force distribution patterns are given in Table 1.

The constants in Table 1 are given when τ ₀ units are s, W _xz units are m, and c _s units are m / s in [Equation 3]. For all three cases, if the pulse width is small to satisfy the condition 2c _s τ ₀ <W _xz , it can be seen that the change of t _max according to τ ₀ satisfies the linear dependence as shown in [Equation 3].

Volume distribution form p b Homogeneous hexahedron 0.552 0.492 Homogeneous spheroid 0.261 0.482 Gaussian spheroid 0.614 0.290

In Equation 3, t _max also depends linearly on W _xz . Therefore, when plotting the change of t _max according to W _xz using the constants of [Equation 3] and [Table 1], it is possible to confirm the accuracy of the constant values of [Equation 3] and [Table 2]. have.

In Fig. 6, the change of t _max obtained from the uniform hexahedron, uniform ellipsoid, and Gaussian ellipsoidal volume distribution is shown as a function of W _xz . In FIG. 6, the dotted line represents the change of t _max linearly approximated by [Equation 3], and well represents the change of the rise time according to W _xz in the region of 2c _s τ ₀ <W _xz .

An analytical approximation of the displacement generated by the circular focus transducer with volumetric (acoustic radiative force) at the focal point was derived by Sarvazyan et al., And the result is shown in [Equation 4].

Where a and d are the radius and the radius of curvature (focal distance) of the single-element transducer, respectively, and I ₀ is the initial shear intensity of the beam axial ultrasound, γ (gamma) = η _s / ρ, which is the dynamic shear viscosity. , D = 2cd / ωa ² (ω is the angular frequency).

Equation 4 is applied when the pulse width? ₀ of the ultrasonic beam is smaller than the time required for the transverse wave to pass in the transverse direction of the beam, and the beam width in the axial direction in the focal region is much larger than that in the transverse direction.

The displacement at the focal point initially increases with time to reach the maximum (see Fig. 2), and the time to reach the maximum displacement from Equation 4 is given by Equation 5 below.

On the other hand, the acoustic radiation force at the focal plane by the circular focusing probe has a Gaussian distribution and is given by Equation 6 below.

이다. At this time, Is the magnitude of the volume force at the focus,

to be.

이므로, 상기 [수학식 6]은 하기 [수학식 7]과 같이 나타낼 수 있다.Also, a width of -20 dB in the transverse direction at the focal plane

Therefore, Equation 6 may be expressed as Equation 7 below.

Comparing [Equation 7] to [Equation 3], there is no term indicating a dependency on the pulse width τ ₀ of t _max shown in [Equation 3], but the W _xz / c _s dependency is completely consistent. This suggests that the theoretical derivation [Equation 7] is an expression of t _{max at} the limit of τ ₀ → 0, and shows that Equation 3 obtained through numerical calculation is an approximation of a proper function form.

If we plot t _max as a function of τ ₀ in Equation 3, the vertical intercept is the rise time bW _xz / c _s at the limit τ ₀ → 0, so we can estimate the shear wave velocity of the tissue using this intercept value.

For this purpose, it is necessary to calibrate the measuring equipment since acoustic transmissivity imaging has to calculate the shear wave velocity by measuring the displacement-time curve at only one point (focus).

We now describe how to calibrate the measurement instrument by determining the unknown bW _xz by measuring t _max as a function of the pulse width τ _{0 on} a test phantom sample whose shear wave velocity is known as c _s = c ₀ .

As shown in Equation 3, t _max increases linearly with τ ₀ under the condition of 2c _s τ ₀ <W _xz and W _{xz ≪W y} . Therefore, the measurement data tends to be as shown in FIG. 7, and t _max is linearly approximated as shown in Equation 8 below.

In this case, since the vertical axis intercept is β, β = (bW _xz ) / C ₀ compared to Equation 3] _. Since the unknown bW _xz is given by Equation 9 below and the shear wave velocity is known as c _s = c ₀ bW _xz The value can be calculated so that the measuring instrument is calibrated.

Since W _xz of the volume force distribution is difficult to measure directly, a new system parameter K is introduced in combination with an unknown value b, which is an inherent parameter of an ultrasound imaging apparatus for measuring transverse elasticity of human tissue.

If the vertical parameter intercept β is obtained by measuring t _max with an ultrasonic imaging device of which the inherent parameter K is known, the transverse velocity may be given by Equation 10 below.

Therefore, an ultrasound imaging apparatus (i.e., a data acquisition system) is used to transmit ultrasound waves for generating acoustic radiation force to any human tissue, and accordingly, displacement of the tissue (specifically, generated in the transmission direction at the depth of focus of the ultrasound waves). Displacement), which is the intercept of the maximum displacement arrival time in the graph of the ultrasound parameter T of the ultrasound imaging apparatus and the ultrasound transmission time τ-maximum displacement arrival time t _{max as} shown in FIG. If β is found, the sound velocity c _s of the shear wave can be quantitatively obtained.

To this end, in the present invention, as shown in Fig. 8, the transmission time lengths of the ultrasonic sound fields for generating acoustic radiation in the tissues of the human body to be diagnosed are increased to τ ₁ , τ ₂ , and τ ₃ , respectively. Measure the time t ₁ , t ₂ , t _{3 at} which the tissue reaches its maximum displacement.

As described above, the transmission time length τ of the ultrasonic sound field transmitted from the elastic medium and the time t _max at which the medium is pushed out reach the maximum displacement have a linear proportional relationship as described above. As shown in Fig. 9, approximation with a first straight line yields a function as shown in Equation 11, which is an approximation equation corresponding to Equation 8 described above.

Therefore, when [Equation 11] is obtained in this manner, β, which is an intercept (ie, y-axis intercept) of the maximum displacement arrival time t _max for the corresponding human tissue can be obtained arithmetically.

In addition, as described above, when β is obtained, the transverse sound velocity c _s of the tissue can be quantitatively obtained by Equation 10, and when c _s is obtained, the transverse of the tissue is determined by Equation 2. The elastic modulus can be obtained quantitatively.

However, K is an inherent parameter depending on the ultrasonic imaging apparatus, which is a system for obtaining image data as described above, which may be a previously known value, but is known in advance as a sound velocity of a transverse wave generated by acoustic radiation force, as described below. Can be obtained using a test phantom.

On the other hand, in order to quantitatively obtain the sound velocity of the shear wave by the above-described method, it is necessary to acquire data on the displacement of the tissue generated by the acoustic radiation force. In this embodiment, the following method is applied. .

First, as described above, the movement of the soft tissue of the human body, which is pushed by the acoustic radiation generated when transmitting high-power ultrasonic waves for hundreds of microseconds, increases the distance the softer tissue moves.

Therefore, after applying the sound pressure of the ultrasound, measuring the maximum distance moved by the human tissue or measuring the time it takes to move to the maximum distance can determine the tightness of the tissue, when the transmission of the high-power ultrasound stops the tissue again It will return to the position.

The movement of human tissue by the acoustic radiation force is generated at a distance of several tens of micrometers for several seconds. In order to observe the movement, the present embodiment obtains and stores an image at high speed for a short time.

The elastic properties of the tissue can be known by observing the displacement of the image point to be observed with time using the stored data.

10 is a graph for explaining the above-described process, a graph at the top of the graph showing the displacement of the tissue by the acoustic radiation force, the graph in the middle shows the transmission time of the ultrasonic sound field for generating the acoustic radiation force The graph at the bottom is a graph describing the process of obtaining an image of the displacement of the tissue.

In detail, first, when a high-power ultrasonic sound field for the generation of acoustic radiation is transmitted, the tissue starts to push, and when the transmission is stopped, the tissue returns to its original position. In this case, the transmission for the movement tracking is performed for a short time interval. Iteratively obtains an image for tracking the movement.

At this time, in order to observe the displacement of the medium (that is, a test phantom or any ultrasonic elastic phantom such as human tissue to be diagnosed) according to the progress of the acoustic radiation or the transverse wave, an image must be obtained at a very high speed. Since transmission is focused on each scan line, an image of about 30 frames per second can be obtained, which is not suitable for observing a rapidly changing medium movement.

Therefore, in the present embodiment, ultrasonic pulses pass through all regions to be obtained by transmitting pulse plane waves having a very short time length without focusing, and the received reflected signals are collected at once to provide dynamic reception focusing for all image points. By doing this, it is possible to obtain the image of the whole area in one transmission.

Such a method has a disadvantage in that the resolution is lowered compared to the conventional medical ultrasound image that focuses on transmission, but it has an advantage that image data can be acquired at a speed of about 10,000 frames per second to obtain an image having a depth of up to 50 mm. (Actually, the image data only needs to obtain a period of several tens of microseconds in which the medium is moved by the acoustic radiation as described above.)

To this end, in the present embodiment, an ultrasonic imaging apparatus for obtaining image data at high speed is manufactured, and a detailed configuration thereof is shown in FIG.

The ultrasound imaging apparatus includes an ultrasound imaging module 10 for receiving ultrasound transmission and reflection signals through a linear transducer 40, an interface 20 connected to the ultrasound imaging module 10, and storing image data, and the It is configured to include a PC 30 for signal processing using the image data stored in the interface 20.

The ultrasound imaging module 10 has been modified to transmit a plane wave of short pulse length in the conventional medical ultrasound imaging apparatus (GE, LOGIQ P6), and the interface 20 can store highway image data of 9800 frames per second. It was configured by hardware.

The data acquisition system configured as described above amplifies a signal received at each element of the transducer 40 and transmits the analog signal to the interface 20 as an analog signal. It is configured to.

The data stored for about 30 ms is transferred to the PC 30 after the data acquisition is completed, and the signal processing is performed. In this embodiment, the signal processing is performed using MATLAB as an example.

In addition, the motion calculation of the ultrasonic elastic phantom was calculated by comparing the position of each image point in the image of the first reference frame with no motion and the image of the frame with motion, which is perpendicular to the transducer 40. Since the movement in the scan line direction is the largest, the displacement was calculated using the one-dimensional auto-correlation calculation method in the scan line direction.

A method of quantitatively measuring the transverse elasticity of human tissue according to the present invention using the ultrasonic imaging apparatus configured as described above will be described in detail.

First, while varying the transmission time length τ of the ultrasonic transmitting sound field for generating acoustic radiation for the test ultrasonic elastic phantom which knows the sound velocity characteristics of the shear wave in advance, the test ultrasonic elastic phantom reaches the maximum displacement. By measuring the time t _max , a graph as shown in Fig. 8 is obtained.

To this end, in the present embodiment, the test ultrasonic elastic phantom is manufactured by mixing liquid plastic, a curing agent, and a softener (MF Manufacturing Co., USA) as an example, and the ultrasonic reflector is colorless and has a diameter of about 27 μm. Flour was added at 0.5% of the total weight ratio.

In addition, the test ultrasonic elastic phantom was produced to have two kinds of elastic properties by adjusting the weight ratio of the curing agent and softener, each physical property is shown in the following [Table 2].

Kinds Plastic
Hardener [g] Plastic
Softener [g] Scatterer
[g] density
[kg / ㎥] Young's
Modulus [kPa] soft phantom 70 130 One 1060 11.1 hard phantom 140 60 One 960 20.4

From the time-displacement data (ie, a graph as shown in FIG. 8) of the test ultrasonic elastic phantom obtained by the above-described method, a linear linear approximation such as [Equation 8] or [Equation 11] is performed. After the relationship graph of the ultrasonic transmission time (tau)-the maximum displacement arrival time (t _max ) is calculated | required, the y-axis intercept (specifically, the intercept of the maximum displacement arrival time) (beta) of Formula (11) is obtained from this.

In this case, since the sound velocity of the transverse wave is previously known in the case of the test ultrasonic phantom, the inherent parameter K of the ultrasonic imaging apparatus is calculated using Equation 10.

As such, when the value of the inherent parameter K of the ultrasound imaging apparatus is obtained, the velocity and the transverse elasticity value of the shear wave for any human tissue can be quantitatively calculated using the same ultrasound imaging apparatus.

That is, the arbitrary ultrasonic waves are changed for each case while varying the transmission time length τ of the ultrasonic sound field for generating acoustic radiation force in any ultrasonic elastic phantom (i.e., soft tissue of the human body) that does not know the sound velocity of the shear wave. The time t _max at which the elastic phantom reaches the maximum displacement is measured.

From the time-displacement data of the arbitrary ultrasonic elastic phantoms obtained as described above, a linear straight line approximation as shown in [Equation 11] can be obtained, and when the segment β is obtained, the inherent parameter K of the ultrasonic imaging apparatus is known in advance. Therefore, the sound velocity c _s of the shear wave for the arbitrary ultrasonic elastic phantom can be quantitatively calculated using Equation 10.

Furthermore, when the shear wave velocity c _s for the arbitrary ultrasonic elastic phantom is obtained by the above-described method, the transverse elastic modulus μ for the arbitrary ultrasonic phantom can be quantitatively obtained according to Equation (2). Where ρ is the density of any ultrasonic elastic phantom.

As described in detail above, the method of measuring the transverse elasticity of the human tissue according to the present invention generates the acoustic radiation force in the human tissue to be diagnosed while changing the transmission time of the ultrasonic wave, in which case the time of the human tissue is generated. Since the transverse elastic properties of human tissues can be quantitatively measured using the displacement according to the above-described method, the accuracy of the diagnosis result by the ultrasound imaging apparatus can be improved.

In addition, the method of measuring the transverse elasticity characteristics of the human tissue according to the present invention is not a method of directly measuring the sound velocity of the shear wave, but as a method of measuring indirectly using the displacement according to the time of the human tissue as described above, Since there is no need to directly obtain the speed, the ultrasound imaging apparatus can be configured at low cost.

10: ultrasonic imaging module 20: interface
30: PC 40: Transducer

Claims (5)

Ultrasonic waves are transmitted to the first phantom to be diagnosed by using an ultrasonic imaging apparatus while changing the length of the transmission time τ a plurality of times, and in each case, the displacement of the first phantom in the transmission direction is maximized at the depth of focus of the ultrasonic waves. A first step of measuring the maximum displacement time (t _max ) to be; And
The correlation between the transmission time (tau) and the maximum displacement time (t _max ) obtained from the measurement result of the first step, and the inherent parameter (K) of the ultrasound imaging apparatus known in advance, And a second step of obtaining the transverse elastic properties. The method of claim 1,
The second step may include: a step 2-1 of approximating the maximum displacement time t _{max as} a first function of a transmission time τ from the measurement result of the first step; and
And a step 2-2 of obtaining a lateral elasticity characteristic of the first phantom using the intercept β of the maximum displacement time t _max with respect to the linear function and the inherent parameter K of the ultrasound imaging apparatus. Method for measuring the lateral elastic properties of human tissue, characterized in that. The method of claim 2,
The lateral elasticity characteristic of the first phantom is the sound velocity (c _s ) of the transverse wave generated in the transverse direction which is a direction perpendicular to the transmission direction of the ultrasonic wave by vibrating the first phantom by the transmitted ultrasonic waves,
The second step 2-2 is to determine the sound velocity (c _s ) of the shear wave according to the formula [1] below.
[Equation 1]

The method of claim 3,
Step 2-2 further includes the step of obtaining a transverse elastic modulus (μ), which is a lateral elastic modulus of the first phantom, from the sound velocity (c _s ) of the shear wave according to Equation 2 below. A method of measuring the lateral elasticity of human tissue.
[Equation 2]

5. The method according to any one of claims 1 to 4,
The unique parameter K of the ultrasound imaging apparatus is
By vibration by the ultrasonic transmit acoustic velocity of transverse wave generated by the transmission in a direction normal to the direction of the ultrasonic wave is also (c _s0) a, while a plurality of times to change the length of the transmission time (τ ₀₎ with the second phantom already-known ultrasound In each case, the maximum displacement time t _max0 at which the displacement of the second phantom is maximized in the transmission direction at the depth of focus of the ultrasonic wave is measured, and the maximum displacement time t _max0 is calculated from the measurement result. τ ₀ ) after approximating the first function, the transverse elastic properties of the human tissue, characterized in that obtained by applying the intercept (β ₀ ) of the maximum displacement time (t _max0 ) for the first function to the following [Equation 3] How to measure.
[Equation 3]

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