JPWO2010147055A1 - Ultrasonic imaging apparatus and ultrasonic imaging method - Google Patents

Abstract

Measures intracardiac pressure with non-invasive, high accuracy, and high time resolution. The pressure is calculated from the attenuation curve and resonance curve of the signal reflected from the ultrasound contrast agent. In addition, a multi-component droplet type contrast agent is injected into the body, the composition is set so that the pressure range to be detected becomes gas-liquid equilibrium, and the bubble diameter depends on the pressure, increasing the pressure to the attenuation curve and resonance curve. Reflect and realize highly accurate pressure measurement.

Description

  The present invention relates to an ultrasonic imaging apparatus and method for medical use, and more particularly to an ultrasonic imaging apparatus and method for measuring intracardiac pressure with noninvasive, high accuracy, and high time resolution.

  Heart disease is a disease that affects all people in terms of dysfunction associated with aging, and is one of the three leading causes of death in many developed countries. In such a long-term medical treatment of a disease with many patients, there are wide phases from preventive medical care to advanced medical care. However, as a technical need common to all phases, there is a small burden on the patient. If there is an examination technique with a small burden on the patient, early diagnosis can be made earlier, follow-up can be made more finely, and the quality of medical care can be improved.

  Blood pressure measurement is the basis of testing for heart disease. The main blood pressure measurement methods currently used are compression tests that measure the compression pressure at which the upper arm and other parts are compressed stepwise with a cuff in the preventive medicine, and the pulsating sound disappears. Is a catheter test in which the pressure at the end of the tube left outside the body after insertion is measured.

  Moreover, although it is a research stage, the method of measuring blood pressure non-invasively with an ultrasonic imaging device is reported (patent document 1). This method is a method in which a contrast agent irradiated with ultrasonic waves in a liquid discovers a phenomenon that a subharmonic wave having a signal intensity that linearly reflects the pressure of the liquid is discovered and used. As a procedure, after intravenously injecting an ultrasonic contrast agent, the heart is imaged by an ultrasonic diagnostic apparatus, the signal intensity (dB) of the subharmonic wave is detected, and the signal intensity and pressure of the subharmonic wave that have been examined in advance. The received signal is converted into blood pressure using the relationship.

US Pat. No. 6,302,845

  However, the above-described conventional technology cannot achieve both noninvasiveness, measurement accuracy, and time resolution. Specifically, since the measurement value of the compression test is not the intracardiac pressure, the accuracy is insufficient to be regarded as the intracardiac pressure, and it cannot be used as a definitive cardiac test. Catheter examination is invasive and burdensome to the patient.

  Although the method of patent document 1 is a method of measuring an intracardiac pressure non-invasively, there exists a problem in accuracy and time resolution. Specifically, first, the time resolution and the blood pressure accuracy are not compatible. In this method, the pressure is calculated from the subharmonic wave, but generally the subharmonic wave cannot be obtained at a high S / N ratio. Therefore, the SN ratio is increased by measuring multiple heartbeats and taking the average value of the received signals having the same time phase, or by setting the time interval and taking the average value of the received signals within the time interval. . In other words, instead of improving accuracy, time resolution is sacrificed. However, when measuring the intracardiac pressure as a medical practice, a time resolution of 0.1 seconds or more is necessary, and the time-phase average processing of multiple heartbeats has the problem of excluding arrhythmia patients from the examination target, which is not preferable. . This is because the value used for diagnosis is a blood pressure that specifies the timing of the region and the pulsation cycle, such as left ventricular end-diastolic pressure. Also, time resolution is necessary to prevent arrhythmic patients from being examined. Non-healthy people are important medically important test subjects, and they often have arrhythmias. Next, there is a problem that accuracy cannot be obtained in a low pressure range. The low pressure range (20 to 60 mmHg) is a range of left ventricular end-diastolic pressure that plays an important role in the diagnosis of heart disease, and requires an accuracy of ± 5 mmHg or more. As described above, the method disclosed in Patent Literature 1 is insufficient as a medical device because the accuracy required in a range of values important for diagnosis cannot be obtained.

  The present invention provides a medical imaging apparatus that measures intracardiac pressure with noninvasive, high accuracy, and high time resolution.

  The ultrasonic imaging apparatus according to the present invention includes an ultrasonic probe that transmits / receives ultrasonic waves to / from an imaging target injected with an ultrasonic contrast agent, and an ultrasonic signal from the imaging target received by the ultrasonic probe. A signal processing unit for processing, and display means for displaying a processing result of the signal processing unit. The signal processing unit calculates the frequency distribution of the received ultrasonic signal, and is calculated by the frequency analysis unit. A frequency distribution that is one or more of a value at one or more specified frequencies, a frequency indicating a maximum value, a maximum value, and a half-value width of the frequency dependence of the attenuation rate of the received ultrasonic signal from the frequency distribution. A frequency distribution feature quantity detection unit for detecting the feature quantity; and a frequency distribution feature quantity-pressure conversion section for converting the frequency distribution feature quantity to pressure, and the display means is calculated by the frequency distribution feature quantity-pressure conversion section. Displays the pressure value.

  The ultrasonic imaging method of the present invention includes a contrast agent injection step for injecting an ultrasonic contrast agent into an imaging target, an ultrasonic transmission step for transmitting an ultrasonic signal to the imaging target, and a reflection from the imaging target. A frequency analysis step for receiving an ultrasonic signal and calculating a frequency distribution of the received ultrasonic signal, and one or more specified frequencies of the frequency dependence of the attenuation rate of the received ultrasonic signal from the calculated frequency distribution A frequency distribution feature quantity step for detecting a frequency distribution feature quantity that is at least one of a value at, a frequency indicating a maximum value, a maximum value, and a half width, and a frequency distribution for converting the detected frequency distribution feature quantity into pressure And a feature amount-pressure conversion step.

The ultrasonic contrast agent is a multi-component droplet type contrast agent containing a plurality of liquid components, and the temperature of a part where pressure is detected is T _c , the upper limit of the pressure to be detected is P _u , and the lower limit is P _l . to time, at least one liquid component vapor pressure at temperature T _c is greater than or equal to P _u, the other at least one is not more than the vapor pressure P _l at the temperature T _c, a plurality of liquid components P _l or more, which is preferably mixed in a molar fraction of vapor-liquid equilibrium is established at pressures P _u. Further, at this time, the ultrasonic wave transmission is a two-stage transmission in which an ultrasonic wave for vaporization for vaporizing the multi-component liquid droplet type contrast agent is performed and then an ultrasonic wave for pressure detection is transmitted. .

  According to the present invention, the pressure in the body can be measured with non-invasive, high accuracy, and high time resolution. More specifically, while the conventional technique has signal strength at one specified frequency, the present invention uses physical quantities at a plurality of frequencies and uses a large number of sample points, so that high accuracy can be obtained. In addition, in order to extract the subharmonic which is a non-linear component in the conventional example, it is necessary to transmit twice for each reception. However, the present invention does not need that, and even simple calculation is twice as high as the conventional example. Time resolution is obtained.

  Furthermore, in one embodiment of the present invention, a multi-component droplet type contrast agent containing two or more liquid components is injected into the body, and the temperature is increased by irradiating the focused site with ultrasonic waves. The liquid contained in the agent is vaporized to establish a vapor-liquid equilibrium state. In the vapor-liquid equilibrium state, the volume of the gas phase changes depending on the pressure, but this volume change appears as a change in the bubble diameter.As a result, the behavior of the bubbles contained in the received signal is enhanced by the influence of pressure and with high accuracy. The pressure can be detected.

1 is a block diagram showing a configuration example of an ultrasonic imaging apparatus according to the present invention. The flowchart showing the process of the ultrasonic imaging method by this invention. The flowchart showing the detail of the process which calculates a pressure from a received signal. The flowchart showing the detail of the process which calculates the frequency distribution of attenuation amount. Explanatory drawing of the process which transmits and receives an ultrasonic wave to a biological body. Explanatory drawing of the process which calculates the frequency distribution of attenuation amount. The figure showing distribution of the bubble diameter of an ultrasonic contrast agent. The figure showing the example of a frequency distribution feature-value, and the example of the attenuation | damping curve at the time of changing a pressure. The figure explaining the process which transmits and receives an ultrasonic wave to a biological body. The flowchart of the process which calculates the frequency distribution of attenuation amount. The conceptual diagram which shows the structure of a typical ultrasonic contrast agent. The conceptual diagram of the vapor-liquid equilibrium curve in the constant temperature of a two-component system. The figure which shows the example of the gas-liquid equilibrium music in the constant temperature of a 2 component system. 1 is a block diagram showing a configuration example of an ultrasonic imaging apparatus according to the present invention. The flowchart which shows the process of the Example which displays hardness. The figure explaining the method of detecting the amount of movements in the case of calculating the hardness of a blood vessel wall as an example. Explanatory drawing of the process which transmits and receives an ultrasonic wave to the heart. The flowchart which shows the flow of the whole process which displays the shape and pressure of a measurement site | part. The figure which shows an example of a display.

  FIG. 1 is a block diagram showing a configuration example of an ultrasonic imaging apparatus according to the present invention. The ultrasonic probe 1 is composed of a plurality of elements. The apparatus main body 2 includes a transmission beamformer 3, an amplification unit 4, a reception beamformer 5, a signal processing unit 6, a memory 7, a display unit 8, an input unit 9, and a control unit 10. The signal processing unit 6 includes a frequency analysis unit 61 that calculates the frequency distribution of the received signal, a frequency distribution feature amount detection unit 62 that detects a frequency distribution feature amount from the frequency distribution calculated by the frequency analysis unit 61, and a frequency distribution feature amount. A frequency distribution feature value-pressure conversion unit 63 that converts pressure is provided. The frequency distribution feature quantity detected by the frequency distribution feature quantity detection unit 62 is one of a value at one or more specified frequencies, a frequency indicating a maximum value, a maximum value, and a half-value width of the frequency dependency of the attenuation rate. That's it. The frequency distribution feature value-pressure conversion unit 63 includes a frequency distribution feature value-pressure conversion knowledge storage unit 631 that stores frequency distribution feature value-pressure conversion knowledge, which is knowledge for converting the frequency distribution feature value into pressure. It should be noted that configurations such as normal B-mode display and Doppler display are not shown for simplicity.

  An ultrasonic contrast agent is injected by a user into a living body part to be imaged, and an ultrasonic pulse generated by the transmission beamformer 3 is transmitted from the ultrasonic probe 1 to the living body, from the living body part including the ultrasonic contrast agent. The ultrasonic probe 1 receives the reflected ultrasonic waves. The received signal is input to the amplifying means 4 and amplified, and the receiving beamformer 5 performs phasing addition. The signal processing unit 6 receives the reception signal output from the reception beamformer 5 and performs imaging and pressure calculation. The created image and the calculated pressure are stored in the memory 7, read out and interpolated, and displayed on the display unit 8. These processes are controlled by the control means 10.

  In the signal processing unit 6, after the frequency analysis unit 61 calculates the frequency distribution of the received signal, the frequency distribution feature quantity detection unit 62 detects the frequency distribution feature quantity from this frequency distribution, and the frequency distribution feature quantity-pressure conversion knowledge is obtained. The frequency distribution feature value-pressure conversion unit 63 converts the frequency distribution feature value into pressure.

  2 to 4 are process flowcharts showing an example of the process of the ultrasonic imaging method of the present invention. In the present invention, as shown in FIG. 2, the user first injects an ultrasound contrast agent into the imaging target (S11). Thereafter, an ultrasonic signal is transmitted from the ultrasonic probe 1 to the object to be imaged (S12), the amplifying means receives the ultrasonic signal from the object to be imaged, and performs phasing addition with a reception beamformer. The processing unit 6 performs processing to calculate the pressure (S13). The calculated pressure is displayed on the display means (S14).

  Here, the ultrasound contrast agent used in the present invention is a microbubble having a diameter of several tens of μm or less called a microbubble or nanobubble, and may have a shell composed of a biodegradable polymer.

FIG. 3 shows details of processing of the signal processing unit 6 in step 13 of FIG. The signal processing unit 6 calculates the frequency distribution of the attenuation amount based on the received signals having different reception depths (S21), and detects the frequency distribution feature quantity from the frequency distribution (S22). After that, the frequency distribution feature value is converted into the pressure p ₀ using the frequency distribution feature value-pressure conversion knowledge shown in the following equation (1), which is an ideal attenuation frequency distribution without pressure measurement noise. (S23). As a conversion method, for example, a frequency distribution feature amount is input, and a pressure that is a parameter is obtained by a least square method using frequency distribution feature amount-pressure conversion knowledge as a regression curve.

Here, σ _e is a scattering cross section, ω is a frequency which is a variable of the formula (1), normalized by the resonance frequency ω ₀ , and expressed as Ω. a _e is the equilibrium value of the bubble diameter, ρ _L is the density of the surrounding fluid, κ is the polytropic index, G _S is the rigidity of the bubble shell, d _Se is the shell thickness defined by equation (3), η _S Represents the shear viscosity of the bubble shell and is determined by the physical properties of the contrast agent and fluid such as blood. These are constants in the equation (1), and are examined in advance and stored in the apparatus. p ₀ is the hydrostatic pressure, the only parameter in (1), the pressure to be detected by the apparatus and method of the present invention.

FIG. 4 is a diagram showing details of the process of calculating the frequency distribution of the attenuation amount in step 21 of FIG. First, received signals Sig _i (t) having different received depths and received depths z _i {i = 1, 2,..., N} are input (S31). Next, the frequency distribution | Sig _i (f) | = | F [Sig _i (t)] | of the received signal Sig _i (t) {i = 1, 2,..., N} is calculated (S32). Finally, the frequency distribution of attenuation from the frequency distribution | Sig _i (f) |
σ (f) ＝ | Sig _i (f) | ／ | Sig _j (f) |
Is calculated (S33). Here, F [•] represents a Fourier transform.

Next, details of the imaging method by the ultrasonic imaging apparatus of the present invention will be described with reference to FIGS. FIG. 5 is a diagram illustrating a state in which the ultrasonic probe transmits and receives ultrasonic waves to a living body. When ultrasonic waves are transmitted from the ultrasonic probe 1 to the living body 72, the signals 75 and 76 reflected from the ultrasonic contrast agent 74 included in the organ 73 are received. The reception depth of the signal 75 is z ₁ and the reception depth of the signal 76 is z ₂ . The organ 73 may be an organ containing a fluid whose reflection luminance is significantly lower than that of an ultrasound contrast agent, such as a blood vessel or a heart chamber of the heart, or may be an organ composed of a soft tissue having a reflection luminance of a level that can be visually recognized, such as a liver. . In the latter case, the signal processing unit receives signals similar to the reception signals 75 and 76 when there is no ultrasonic contrast agent so that the frequency analysis unit can process only attenuation by the ultrasonic contrast agent, What is necessary is just to perform the process which subtracts the attenuation amount resulting from other than an ultrasonic contrast agent. It should be noted that the attenuation caused by other than the ultrasonic contrast agent is generally an amount proportional to the frequency, and can be obtained by, for example, transmitting / receiving / analyzing ultrasonic waves before injection of the ultrasonic contrast agent.

FIG. 6 is a diagram for explaining the processing of step 21 in FIG. 3 for calculating the frequency distribution of the attenuation amount. A distribution 81 indicates a signal Sig _i (f) = F [Sig _i (t)] obtained by performing a Fourier transform on the reception signal Sig _i (t) from the reception depth z _i , and a distribution 82 indicates the signal intensity at the reception depth z _j. Frequency distribution. In step 32 of FIG. 4, such frequency distributions 81 and 82 are calculated at a plurality of reception depths. However, in the received signal Sig _i (t) from the reception depth z _i , c is a constant representing the speed of sound, dt is a time including the number of samples necessary for frequency calculation, and the time is an ultrasonic wave sent to the conversion imaging target. Assume that the waved time is t = 0, and the signal is received at a time corresponding to the reception depth z _i (from t = z _i / c-dt to t = z _i / c + dt).

  The distribution shown in FIG. 6B is a frequency distribution σ (f) of attenuation expressed by the following equation, which is obtained by taking a ratio of frequency distributions 81 and 82 of signal intensity. In step 33 of FIG. 4, the frequency distribution σ (f) of this attenuation is calculated. However, in FIG. 6, it is shown in dB.

σ (f) ＝ | Sig _i (f) | ／ | Sig _j (f) |
Here, the frequency distribution feature amount-pressure conversion knowledge represented by Expression (1) will be described. An ultrasound contrast agent is a microbubble around several μm that is injected into the body, and when irradiated with ultrasound, the outside of the bubble is a liquid and the inside is a gas, and the density difference is large. Reflects ultrasound signals. Usually, the signal intensity of the fundamental wave component and the high frequency component of the signal reflected from the ultrasound contrast agent is imaged. In this way, in normal usage, the presence of bubbles is used to visualize the boundaries of blood vessel walls and liver tumors. However, in the present invention, the objective is to extract information beyond the presence of bubbles from the mechanical behavior of the bubbles, and an overview of the properties of the equations describing the bubble motion. Bubble motion is described by the following Church equation, which is the equation of motion for bubble radius a (t).

Where a ₁ is the inner diameter of the bubble, a ₂ is the outer diameter of the bubble, a _1e is the equilibrium value of the inner diameter, a _2e is the equilibrium value of the outer diameter, ρ _L is the density of the surrounding fluid, and ρ _S is the bubble shell Density, η _L is the shear viscosity of the surrounding fluid, η _S is the shear viscosity of the bubble shell, V _S is the volume of the bubble core, p _∞ (t) is the pressure at a location sufficiently away from the bubble, and p ₀ is the hydrostatic pressure The pressure to be detected.

There are various equations describing the behavior of bubbles, depending on the assumptions made, but in order to increase the remaining time in the living body, since bubbles with shells have been the mainstream since 2000, the equations for bubbles with shells have been used. The Church equation is used. Here, the bubble diameter displacement amount x (t) is made a variable by approximation of equation (3) and variable conversion shown in equation (4). a is the bubble diameter and a _e is the equilibrium value of the bubble diameter.

Further, when linearization is performed to obtain Equation (5), Equation (6) is obtained as a solution.

Here, p _i (t) is a pressure change due to irradiation of an ultrasonic pulse.

The frequency distribution of the attenuation rate can be obtained from Expression (7) to Expression (8) of the scattering cross section.

  In the formula (8), n (a) is a bubble diameter distribution. For example, the bubble diameter has a distribution as shown in FIG. 7, but in the above description relating to step 33 in FIG. 4, the distribution width is 0, that is, the bubbles inside the imaging target are all assumed to have the same value. did. Actually, it is necessary to reflect n (a). For example, the distribution of the bubble diameter before injection into the body may be measured and used.

Of the above results, those necessary for qualitative understanding are summarized below.

That is, the shape of the resonance curve (6) is substantially determined by the resonance frequency ω ₀ and the dispersion magnitude δ. The shapes of the attenuation curves (7) and (8) are also substantially determined by ω ₀ and δ and the frequency distribution p _i (ω) of the ultrasonic pulse. Of the right side of the equation (9) representing ω ₀ and δ, values other than p ₀ are fixed by the structure (physical property constant) of the ultrasound contrast agent. Only the shell stiffness G _s and the bubble equilibrium diameter a _e are substantially variable. The shape of the resonance curve (6) and the attenuation curve (7) (8), that is, the sensing power of the pressure p ₀ can be adjusted by these two values, the bubble shell rigidity G _s is small, and the bubble equilibrium diameter a _e is It can be seen that a small ultrasonic contrast agent is suitable for pressure detection.

For reference, Table 1 shows physical property constants of commercially available ultrasonic contrast agents.

FIG. 8 shows an example of a frequency distribution feature amount and an example of an attenuation curve when the bubble has an equilibrium diameter a _e constant and the pressure is changed to three values. The decay curve is significantly deformed by pressure. As frequency distribution features for detecting this deformation, in FIG. 8, the resonance frequency ω ₀ , the dispersion δ ⁻¹ , the attenuation intensity at the specified frequency ω _{1, and} the attenuation intensity at the two specified frequencies ω ₁ and ω ₂ are shown. The inclination was shown.

All the frequency distribution feature values described above are input to the frequency distribution feature value-pressure conversion knowledge of the equation (1) to give the pressure p ₀ as an output. These frequency distribution feature quantities can be simultaneously input to equation (1) even if they are different physical quantities such as resonance frequency and variance, and the detection accuracy of p ₀ can be improved as the number of input values increases. .

More specifically, when the attenuation intensity σ _i (i = 1, 2,..., N) at a predetermined frequency ω _i is obtained, the pressure p ₀ can be obtained as a regression problem with (7) as a regression equation. . That is, by the least square method
J = Σ _{i = 1 to n} | σ _e (ω _i ; p ₀ ) −σ _i | ^2, which minimizes p ₀ ,
δJ = ∂ / ∂p ₀ [Σ _{i = 1 to n} | σ _e (ω _i ; p ₀ ) −σ _i | ² ]
= 0
The value of p ₀ that satisfies the above is obtained numerically and used as the pressure detection value. When considering the influence of the distribution n (a _e ) of the radial direction, (8) should be used as a regression equation and σ _e (ω _i ; p ₀ ) above should be replaced with Σ _e (ω _i ; p ₀ ) .

Further, in addition to the attenuation intensity σ _i (i = 1, 2,..., N) at the specified frequency ω _i, the resonance frequency ω ₀ , the dispersion δ ⁻¹ , and the attenuation at the two specified frequencies ω ₁ and ω ₂ . If intensity gradients are obtained, these are converted into attenuation intensity σ _i (i = n + 1, n + 2,..., N + m) at a prescribed frequency ω _i to increase the number of input values of the regression problem, and p ₀ can be obtained. Or, if only one of these is obtained, the regression problem is reduced to one equation. For example, when only the resonance frequency ω ₀ is obtained, the regression problem is reduced to the equation (7′-2), and ω ₀ may be converted to p ₀ by the equation (7′-2). Similarly, when only the variance δ is obtained, the regression problem is reduced to the equation (7'-2) and the equation (7'-3), and the equation (7'-2) and the equation (7'-3) δ can be converted to p ₀ .

  According to the above configuration, it is possible to provide a medical imaging apparatus that measures the pressure in the body with noninvasive, high accuracy, and high time resolution. In particular, since signal strengths at a plurality of frequencies are used and a large number of sample points are used, the S / N ratio is increased and high accuracy is obtained. Further, unlike subharmonic wave extraction, it is not necessary to transmit twice or more in order to detect the pressure at one spatial position, so that high time resolution can be obtained. Furthermore, in the present invention, accuracy can be improved by detecting one or more physical quantities and converting them into blood pressure.

  Next, a second embodiment of the present invention will be described with reference to FIGS. The method of the second embodiment is an example in which the resonance curve shown in Expression (10) is used as the frequency distribution feature amount-pressure conversion knowledge in the processing of step 13 in FIG.

  FIG. 9 is a diagram illustrating a process in which the ultrasonic probe transmits and receives ultrasonic waves to the living body in the present embodiment. When an ultrasonic wave is transmitted from the ultrasonic probe 1 to the living body 72, the signal 75 reflected from the ultrasonic contrast agent 74 included in the living body is received. The organ 73 may be an organ containing a fluid whose reflection luminance is significantly lower than that of an ultrasound contrast agent, such as a blood vessel or a heart chamber of the heart, or may be an organ composed of a soft tissue having a reflection luminance that can be visually recognized, such as a liver. . In the latter case, it is only necessary to perform the process described with reference to FIG. 5 and extract only the attenuation due to the ultrasound contrast agent.

FIG. 10 is a flowchart for explaining the process of calculating the pressure from the received signal by the signal processing unit 6 executed in step 13 of FIG. 2 in this embodiment. A received signal at a certain receiving depth is inputted and its frequency distribution is calculated (S41). Next, from the frequency distribution, a frequency distribution feature quantity that is one or more of the frequency distribution of the intensity of the fundamental wave, the harmonic wave, the subharmonic wave, the difference frequency, the sum frequency, and the value branch is detected ( After that, the frequency distribution feature value is converted into pressure using the following frequency distribution feature value-pressure conversion knowledge which is an ideal resonance curve without measurement noise using the pressure as a parameter (S43).

For example, when the frequency distribution x _B (ω) of the intensity of the fundamental wave is measured, x _B (ω) should be a resonance curve x (ω). Therefore, if the least square method is performed with the measured value x _B (ω) as the input value and (10) as the regression equation, p ₀ is obtained. Similarly, the frequency distribution x _n (ω) of the intensity of the n-th harmonic, the frequency distribution x _{1 / m} (ω) of the intensity of the 1 / m subharmonic, and the frequency distribution p _m−n ( ω), when the frequency distribution pm _{+ n} (ω) of the intensity of the sum frequency is measured, p _i (ω) in the regression equation (10) is set as the nth-order harmonic assumed when there is no contrast agent. _Pn is calculated by the regression equation (10) as the frequency distribution of the intensity of 1 / m subharmonic, difference frequency, and sum frequency. For example, when one or more frequency feature quantities such as a fundamental wave and a harmonic wave are obtained, the number of regression problem inputs is increased as in the first embodiment, and p ₀ is obtained more accurately. It is done.

  According to the configuration as described above, the pressure can be obtained from the signal of the ultrasound contrast agent in the capillary blood vessel or organ in which the region having a uniform structure is small and the attenuation cannot be detected with good S / N.

  Next, a third embodiment of the present invention will be described with reference to FIGS.

In the first and second embodiments, as described using the equation (9), the signal of the ultrasonic contrast agent reflects the pressure p ₀ and the equilibrium diameter a _{e of the} bubble. Bubbles with shells that have been used to increase the remaining time in the body since 1980. In equation (9), if the stiffness G _{s of the} bubble shell is large, the behavior of the bubbles, and thus the resonance curve The decay curve more strongly reflects the bubble equilibrium diameter a _e than the external pressure p ₀ . Therefore, an example in which the pressure detection accuracy is improved by using an ultrasonic contrast agent containing a plurality of components in the liquid state to make the bubble equilibrium diameter a _e agile function of the external pressure p ₀ will be described below. This is a third embodiment.

In the third embodiment, the composition of the ultrasound contrast agent is different, and the relationship between the equilibrium diameter a _{e of the} bubbles and the external pressure p ₀
a _e = f (p ₀ ) (11)
Is used together with the formula (1) or the formula (10). It is assumed that the function f (p ₀ ) is obtained from experiments and stored in the apparatus in advance. However, except for the gas-liquid two-phase coexistence state, it may be theoretically obtained from the van der Waals equation of state and the molar fraction of the constituent components and stored in the apparatus.

  FIG. 11 is a conceptual diagram showing the structure of a typical ultrasonic contrast agent. FIG. 11A shows an ultrasound contrast agent that contains a gas inside, does not have a shell, and has a bubble diameter d of 1 μm or more. Such an ultrasound contrast agent has already been commercialized and is called a first generation contrast agent, and has a short remaining time in a living body. FIG. 11B shows an ultrasound contrast agent that contains a gas inside, has a shell 141, and has a bubble diameter d of 1 μm or more. Such an ultrasound contrast agent has already been commercialized and is called a second generation contrast agent, and has a long remaining time in a living body. The thickness t of the shell 141 is several percent or less of the diameter d, and a polymer having a hydrophilic portion 142 and a hydrophobic portion 143 is a structural unit. FIG. 11C shows an ultrasound contrast agent that contains a liquid inside, has a shell 144, and a bubble diameter d of 1 μm or less. Such an ultrasound contrast agent has not yet been commercialized and is in the research stage. It is called a phase change droplet type contrast agent. The penetrable region such as the tip of the capillary is wide, and is vaporized at an intended site by ultrasonic irradiation and used for imaging. The third embodiment is an example using a phase change droplet type contrast agent containing a plurality of types of liquids.

  The phase change droplet type contrast agent currently being studied includes that of Hitachi, Ltd. Kawabata et al. The physical properties are as follows.

Shell material: Lipid / surfactant emulsion
Droplet diameter: 200nm
Inclusion: PEOB
In the third embodiment using a phase change type contrast agent, a gas-liquid phase change type multi-component contrast agent is injected into the imaging target in step 11 of FIG. Further, in the process of transmitting the ultrasonic signal to the imaging target in step 12, first, the ultrasonic wave for vaporization is transmitted to the imaging target to vaporize the liquid inside the contrast agent, and then the ultrasonic wave for pressure detection is supplied. Transmit to the imaging target. However, when the vapor pressure of the component contained in the contrast agent is lower than the pressure value range to be measured, this step of transmitting the ultrasonic wave for vaporization to the imaging target may be omitted. When transmitting, the spatial position where the ultrasonic wave for vaporization is transmitted is the same as the spatial position where the ultrasonic wave for pressure detection is transmitted or received, or the position where the blood flow is upstream. For example, when detecting the pressure in the left ventricle of the heart, an ultrasonic wave for vaporization may be transmitted into the left ventricle or the left atrium. Further, when detecting the pressure in the liver, an ultrasonic wave for vaporization may be transmitted to the hepatic artery.

Next, using FIG. 12 and FIG. 13, it will be explained that the bubble equilibrium diameter a _e can be made an agile function of the external pressure p ₀ by using an ultrasonic contrast agent containing a plurality of components in a liquid state. FIG. 12A is a one-component isothermal curve, that is, Van der Waals equation of state, and FIG. 12B is a conceptual diagram of a vapor-liquid equilibrium curve at a constant temperature of the two-component system.

  In FIG. 12A, the vertical axis represents pressure, and the horizontal axis represents volume. The change according to the van der Waals equation of state shown at the bottom of the figure is shown by a dotted line. The solid line in the figure is the volume-pressure change followed by the actual fluid. However, the temperature is assumed to be constant. The actual fluid volume-pressure occurs monotonically, unlike Van der Worth's equation of state. The difference from the equation of state is a straight line drawn so that the areas S1 and S2 in the figure are equal. This straight line where the volume changes but the pressure is constant is the gas-liquid phase coexistence in the gas-liquid phase change. In this state, the pressure at this time is called a vapor pressure Pc. On the line where the pressure is lower than the gas-liquid two-phase coexistence state, the gas state is present, and on the larger line is the liquid state. As described above, in the one-component system, if the pressure is changed while keeping the temperature constant, the volume of the vapor pressure Pc changes greatly with a discontinuous jump. In the multi-component system, if the pressure is changed while keeping the temperature constant, The volume always changes continuously. For this reason, the pressure change can be quantified by the volume change.

FIG. 12B shows the volume change in the multi-component gas-liquid phase change. In FIG. 12B, a two-component system is assumed. The vertical axis represents pressure, and the horizontal axis represents the molar fraction of the first component. Pc1 is the vapor pressure of component 1, Pc2 is the vapor pressure of component 2, thick solid lines mean liquid phase lines, and thin solid lines mean gas phase lines. In the region where the pressure is higher than the liquid phase line, it is in the liquid state, in the region where the pressure is lower than the thin solid line, it is in the gas state, and in the middle, the pressure corresponding to the hatched portion is in the gas-liquid two phase coexistence state. In FIG. 12 (a), the pressure in the gas-liquid two-phase coexistence state is determined to be one value Pc, whereas in the two-component system, the hatched portion has a width with respect to the vertical axis P. The pressure in the liquid two-phase coexistence state has a width from the vapor pressure Pc1 of component 1 to the vapor pressure Pc2 of component 2. When the pressure is increased from the gaseous state, when the rate of change is so slow that it can be regarded as a quasi-static process, the molar concentration of component 1, which is a low vapor pressure component, increases through a pressure-molar fraction change like a chain line. , Condensed. The change in the ultrasonic imaging apparatus and method of the present invention can be regarded as a change like a practice arrow. The relationship (11) of the relationship P ₀ = f (a _e ) between the pressure P and the bubble diameter a _e is the van der Waals equation of state (however, the relationship V = πa _e ³ ) except for the gas-liquid two-phase coexistence state (shaded portion). The volume V is converted into a bubble diameter a _e by / 4). It is assumed that the value obtained by experiments is used for the relationship in the gas-liquid two-phase coexistence state (shaded portion).

  The quantitative accuracy of the pressure change becomes higher as the thickness 18 in the pressure direction of the two-component gas-liquid mixed phase is larger. The thickness in the pressure direction of the two-component gas-liquid mixed phase is determined by the physical property values of the components, specifically, the activity coefficient and the vapor pressure. Of course, the closer the molar fraction is to 0.5, the thicker the thickness. That is, utilizing the pressure-volume change in the gas-liquid equilibrium of the multi-component system has the effect of enhancing the pressure dependence in the pressure range intended for measurement.

  The activity coefficient that affects the thickness 18 is a physical property value that represents the degree of interaction of each component in a two-component mixed solution, and it is desirable to use two components having a value of 1 or more as elements of the contrast agent. A specific example (methanol-water system) of FIG. 12 (b) in the case of 1 or more is shown in FIG. 13 (a), and an example of 1 or less (tert butanol-sec butanol system) is shown in FIG. 13 (b). In FIG. 13A, the thickness of the two-component gas-liquid mixed phase in the pressure direction is larger. FIGS. 13 (a) and 13 (b) are examples in which the temperature is 25 degrees, but in FIG. 13 (a), highly accurate pressure measurement is possible in the range of 40 mmHg to 80 mmHg.

When the activity coefficient of the inclusion think that you set as large as possible, as the inclusions of phase shift contrast agent, if compared with _{C 3 F 3, C 3 F} 8, C 5 F 12, SF 6, C 3 F ₃ is good. In addition, adjustment of the molar fraction to around 0.5 is also effective.

If the inclusion component is selected as described above, the bubble equilibrium diameter a _e can be made into an agile function (11) of the external pressure p ₀ . In the third embodiment of the present invention, in the same processing as in the first or second embodiment, p ₀ is obtained by using a formula obtained by substituting (11) for (1) or (10), respectively.
According to the above configuration, the pressure in the body can be detected with high accuracy.

  Next, a fourth embodiment of the present invention will be described with reference to FIGS. This embodiment includes the processing of the first to third embodiments, detects the amount of movement in addition to the pressure at the spatial position where ultrasound is transmitted and received, and the hardness at the spatial position where ultrasound is transmitted and received from the pressure and motion. Is an example of calculating.

  FIG. 14 is a block diagram showing a device configuration of the present embodiment. The signal processing unit 6 of this embodiment includes a motion amount detection unit 64 and a motion amount / pressure-hardness conversion unit 65.

  FIG. 15 is a flowchart showing the flow of processing of this embodiment. The pressure of the imaging target is calculated by the method described in the first to third embodiments (S51). At the same time, the amount of motion of the imaging target is also detected (S52). Any existing method may be used as the method for detecting the amount of motion. For example, a tissue tracking method may be used in which the speckle pattern of the B image is tracked to obtain a movement amount. Thereafter, the hardness is calculated from the pressure and movement (S53), and the hardness is displayed (S54). Hardness is defined as an amount proportional to the pressure and inversely proportional to the amount of movement.

  FIG. 16 is a diagram illustrating a method for detecting the amount of motion, taking as an example the case of calculating the hardness of a blood vessel wall in a blood vessel. FIG. 16A is an overall view of a blood vessel. The method of the present invention first images the blood vessel 101. The imaging section may be either the radial section 102 or the axial section 103. FIGS. 16B and 16C show the imaging results when the radial section 102 is taken, and FIGS. 16D and 16E show the imaging results when the axial section 103 is taken. FIG. 16B and FIG. 16C, FIG. 16D and FIG. 16E show imaging results at adjacent times.

  16B and 16C will be described. FIG. 16B shows the imaging result at time t, and FIG. 16C shows the imaging result at time t + dt. It is assumed that the pressure at the position 104 at time t and the position 104 'at time t + dt is calculated. At this time, the micro area 105 is set in the imaging result at t = t, the micro area 105 ′ determined to have the same shape is detected at t = t + dt, and the center of gravity of the micro areas 105 and 105 ′ is moved. Let the amount 106 be a movement. Further, when the accuracy of the motion amount may be low, for example, a motion vector may be used in which a half of the difference between the blood vessel diameters d1 and d2 is large and the normal direction of the blood vessel wall is a direction. Similarly, the movement amount 106 can be calculated in the axial section 103 shown in FIGS. 16D and 16E.

  According to the above configuration, the hardness of the living body can be detected non-invasively with high accuracy.

  Next, an example in which the processing of the first to fourth embodiments is applied to the heart will be described with reference to FIGS. 17 and 18.

  FIG. 17 is a diagram illustrating a process in which the ultrasound probe transmits and receives ultrasound to the heart. Ultrasound is transmitted from the ultrasound probe 1 to the living body 72, and signals 751 and 752 reflected from the ultrasound contrast agent 74 included in the heart 73 are received.

  The overall flow of the processing is shown in FIG. First, the shape of the heart is picked up by an ultrasonic image (S61), and a part for measuring pressure is detected (S62). Next, the pressure of the measurement site is detected by the method of Example 1 to Example 4 (S63), and the shape and pressure of the measurement site are displayed (S64).

An example of display is shown in FIG. FIG. 19A shows an example in which the blood pressure inside the heart is detected for each location and displayed superimposed on the morphological image. FIG. 19B shows an example in which the result of analyzing the heart morphological image and the detected blood pressure is displayed. is there. 19 (a) and 19 (b), 91 is a display screen, 92 is an imaging region, 93 is a morphological image to be imaged, and in FIG. 19 (a), 94 is a pressure detection result, and different detection values are displayed in different colors. It is a displayed example. In FIG. 19B, reference numeral 95 denotes an analysis result display area, in which the horizontal axis indicates the volume V _LV of the left ventricle extracted by morphological image processing, and the vertical axis indicates the left detected by the imaging method of the present invention. An example is shown in which the spatial representative value P _{LV of the} ventricular pressure is used. _{Reference numeral} 96 shown in the graph represents the time change of (V _LV, P _LV ). Multiple loops are drawn due to differences in measurement conditions. 97 is E _MAX which is one of the indices of the heart function by the inclination of the tangent of the plurality of loops, and 98 is V ₀ which is one of the indices of the heart function by the intercept of the tangent. Although FIG. 19 illustrates an example in which the imaging target is the heart, the imaging target is not limited to the heart. Note that FIG. 19 may be a still image or a moving image.

  According to the above configuration, the blood pressure inside the imaging target such as the heart can be detected noninvasively with high accuracy, and a clinically meaningful amount can be presented to the user together with the morphological information. .

1: Ultrasonic probe 2: Device body 3: Transmission beamformer 4: Amplifying means 5: Reception beamformer 6: Signal processing unit 61: Frequency analysis unit 62: Frequency distribution feature amount detection unit 63: Frequency distribution feature amount − Pressure conversion unit 7: memory 8: display means.

Claims (9)

An ultrasound probe that transmits and receives ultrasound to and from an imaging target injected with an ultrasound contrast agent;
A signal processing unit that performs signal processing of an ultrasonic signal from an imaging target received by the ultrasonic probe;
Display means for displaying the processing result of the signal processing unit,
The ultrasonic imaging apparatus, wherein the signal processing unit sets a focal point at an arbitrary spatial position in an imaging target and detects pressure from a signal reflected by a contrast agent located at the focal point. An ultrasound probe that transmits and receives ultrasound to and from an imaging target injected with an ultrasound contrast agent;
A signal processing unit that performs signal processing of an ultrasonic signal from an imaging target received by the ultrasonic probe;
Display means for displaying the processing result of the signal processing unit,
The signal processing unit is a frequency analysis unit that calculates a frequency distribution of the received ultrasonic signal;
From the frequency distribution calculated by the frequency analysis unit, the frequency dependence of the attenuation rate of the received ultrasonic signal is a value at one or more specified frequencies, a frequency indicating a maximum value, a maximum value, and a half-value width. A frequency distribution feature quantity detection unit for detecting a frequency distribution feature quantity, which is one or more of them,
A frequency distribution feature value for converting the frequency distribution feature value into a pressure-pressure conversion unit;
The ultrasonic imaging apparatus, wherein the display means displays a pressure value calculated by the frequency distribution feature value-pressure conversion unit.   3. The ultrasonic imaging apparatus according to claim 2, wherein the frequency distribution feature value-pressure conversion unit calculates a fundamental wave, a harmonic, a subharmonic wave, a difference frequency, and a sum frequency from the frequency distribution calculated by the frequency analysis unit. An ultrasonic imaging apparatus that detects a frequency distribution feature amount that is one or more of a frequency distribution of intensity and a magnitude of a branch of a value.   3. The ultrasonic imaging apparatus according to claim 2, wherein the signal processing unit is a pressure converted by a motion amount detection unit that detects a motion amount of a part where a pressure in the body is detected, and the frequency distribution feature value-pressure conversion unit. And a pressure / motion amount-hardness amount conversion unit for calculating a hardness amount, which is the hardness of a part where the pressure and motion in the body are detected, from the motion amount detected by the motion amount detection unit. Ultrasonic imaging device. A contrast agent injection step of injecting an ultrasound contrast agent into the imaging target;
An ultrasonic wave transmission step for transmitting an ultrasonic signal to the imaging target;
A frequency analysis step of receiving an ultrasonic signal reflected from the imaging target and calculating a frequency distribution of the received ultrasonic signal;
From the calculated frequency distribution, the frequency dependence of the attenuation rate of the received ultrasonic signal is one or more of a value at one or more specified frequencies, a frequency indicating a maximum value, a maximum value, and a half-value width. A frequency distribution feature amount step for detecting a frequency distribution feature amount;
An ultrasonic imaging method comprising: a frequency distribution feature amount-pressure conversion step of converting the detected frequency distribution feature amount into pressure. The ultrasonic imaging method according to claim 5, wherein
The contrast agent injecting step injects a multi-component droplet type contrast agent containing a plurality of liquid components into the body,
In the ultrasonic wave transmission step, ultrasonic wave for pressure detection is transmitted after performing ultrasonic wave for vaporization for vaporizing the multi-component liquid droplet type contrast agent. Method. 7. The ultrasonic imaging method according to claim 6, wherein when the temperature of a part for detecting pressure is T _c , the upper limit of pressure to be detected is P _u , and the lower limit is P _l , at least of the plurality of liquid components. One has a vapor pressure at a temperature T _c equal to or higher than P _u , and at least one other has a vapor pressure at a temperature T _c equal to or lower than P _l , and the plurality of liquid components are equal to or higher than P _l and lower than P _u. An ultrasonic imaging method characterized by being mixed at a molar fraction at which a gas-liquid equilibrium is established at a pressure of 5. The ultrasonic imaging method according to claim 5, further comprising:
Detecting the amount of movement of the part that detected the pressure in the body;
An ultrasonic imaging method comprising: calculating a hardness amount that is a hardness of a part where the pressure and the motion are detected from the converted pressure and the previously detected motion amount. The ultrasonic imaging method according to claim 6, further comprising:
Detecting the amount of movement of the part that detected the pressure in the body;
An ultrasonic imaging method comprising: calculating a hardness amount that is a hardness of a part where the pressure and the motion are detected from the converted pressure and the previously detected motion amount.

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