JP2003175041A - Ultrasound diagnostic apparatus and image processing method - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide an ultrasound imaging system and a motion information image forming method capable of stably providing motion information image presenting displacement and distortion even in the case of biological tissue such as the heart. <P>SOLUTION: Two-dimensional distribution data of velocity concerning a traveling body in a patient is obtained, information of a motion place defining the moving direction of the tissue is stored, and a plurality of tracking points are decided in a tissue area in the two-dimensional distribution data. Positions which the tracking points move to are estimated and tracked in order, and a moving velocity in a moving direction defined by the motion place is obtained based on the two-dimensional distribution data. In addition, by performing prescribed calculation by using the moving velocity in the moving direction at a tracking position at each time phase at two or more tracking points, a single intermediate output value is obtained and intermediate outputs at different positions are weighted and added, thereby the values of these intermediate positions are obtained. By using the values of these intermediate positions, the motion information image is generated and displayed. <P>COPYRIGHT: (C)2003,JPO

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to an ultrasonic diagnostic apparatus and a motion information image generating method for providing information effective for medical diagnosis.

[0002]

2. Description of the Related Art Objectively and quantitatively evaluating the function of a living tissue such as myocardium is very important for diagnosing the tissue. For example, as a quantitative evaluation method for the heart, there have been the following methods conventionally. As a concrete method, MVG defined by a two-dimensional image is used.
(Myocardial Velocity Grad
ent) method. MV defined by 2D image
G includes MVG-S for short-axis use and MVG-L for long-axis (four-chamber image and two-chamber image by apex approach). In the current practical MVG-S / L, the myocardial site and its local motion direction can be determined by manually setting the contour of the endocardium of the myocardium in a two-dimensional image at a certain time phase. , The velocity gradient of the myocardium, that is, the strain rate is obtained for each region of the myocardium (for example, refer to Patent Document 1).

Another example is TDT (Tissue).
By using the e Doppler Tracking method, if two points (preferably endocardium and adventitia) of the myocardium are given in the initial phase,
There is an MVG-M method that automatically tracks the positions of the two points (endocardium and epicardium) in other time phases and obtains MVG in all time phases (for example, refer to Patent Document 2). There is also a method of obtaining strain in real time and displaying it as a value representing tissue deformation (see, for example, Patent Document 3).

[0004] As an indication of the significance of obtaining a motion parameter in a local region inside the myocardium, in the normal myocardium, the intima side mainly contributes to contraction more dominantly (about twice) than the adventitia side. However, it is said that in the case of a disease such as myocardial infarction, the contribution of the intimal side is reduced (see Non-Patent Document 1, for example). Although it is a property that has been known for a long time in animal experiments, it has been said to be the same in humans in recent studies using the following MVG-M and the like.

Non-invasive and quantitative evaluation of the entire function of the myocardium and the local function inside the myocardium is expected to be useful for grasping the degree of the disease and the selection of a treatment method associated therewith, and is important. It is rising.

However, in the conventional MVG method, in order to analyze a plurality of time phases of the cardiac time phase, it is necessary to set the myocardial region in all the time phases, and the time change of MVG-S / L is analyzed. Is difficult to do. In addition, MVG-S / L is
It is difficult to obtain distribution information in a local region inside the myocardium because the motion parameters are obtained by defining the myocardial size such as the contour of the myocardium or the fractional size of the myocardium as an axis.

Further, in the conventional MVG-M method, since the velocity gradient of the M-mode image is displayed, the time analysis can be performed relatively easily. However, since the MVG-M method has only spatially one-dimensional information, two-dimensional distribution information cannot be obtained.

Further, as a common feature of distortion rate calculation such as velocity gradient, spatial differentiation is performed by using velocity information that tends to be spatially unstable due to the influence of speckle noise peculiar to ultrasonic waves. Therefore, there is a problem that it is difficult to obtain stability in a living body due to noise.

Further, in the conventional distortion display method, the distortion rate between two points having a fixed length in time is obtained, and the distortion rate is spatially uniform and is included in a predetermined section. There is an assumption that These assumptions are
In particular, it is not possible to evaluate the short-axis image, and therefore accurate distortion cannot be obtained.

[0010]

[Patent Document 1] Japanese Patent Laid-Open No. 11-155862

[0011]

[Patent Document 2] Japanese Unexamined Patent Publication No. 9-201361

[0012]

[Patent Document 3] Japanese Patent Laid-Open No. 2001-70303

[0013]

[Non-Patent Document 1] “Difference between Endocardial and Epicardial Myocardium of Left Ventricular Wall” J Cardiol 2000,
No. 35, p. 205-218)

[0014]

SUMMARY OF THE INVENTION The present invention has been made in view of the above circumstances, and can provide a motion information image showing displacement or distortion with high stability even in a living tissue such as a heart. An object of the present invention is to provide an ultrasonic imaging apparatus and a motion information image generation method.

[0015]

The present invention takes the following means in order to achieve the above object.

A first aspect of the present invention is to store a plurality of ultrasonic images of a plurality of time phases of a subject, and a plurality of speeds for each of the time phases based on the plurality of ultrasonic images. A velocity distribution image generating means for generating a distribution image, and a tracking point setting means for setting a plurality of tracking points in the tissue region of the subject in an ultrasonic image regarding a predetermined time phase among the plurality of ultrasonic images, In the plurality of ultrasonic images related to the remaining time phases other than the predetermined time phase, based on the velocity distribution image, estimating means for estimating corresponding points corresponding to the plurality of tracking points, in each of the time phases, Signal value determining means for determining signal values at the tracking points and the corresponding points according to expansion and contraction of the tissue region, and motion for generating a motion information image based on the signal values at the tracking points and the corresponding points Information A generation unit, an ultrasonic diagnostic apparatus characterized by comprising a display means for displaying the motion information image.

A second aspect of the present invention is to store a plurality of ultrasonic images relating to a plurality of time phases of the subject and a first means of the tissue of the subject for the plurality of ultrasonic images.
The center of contraction of the
Center of contraction setting means for setting the center of contraction, a first motion velocity distribution image relating to motion along a direction toward the first center of contraction, and the second contraction based on the plurality of ultrasonic images. A velocity distribution image generating means for generating, for each time phase, a second motion velocity distribution image relating to the motion along the direction toward the center; and an ultrasonic image relating to a predetermined time phase among the plurality of ultrasonic images. , Tracking point setting means for setting a plurality of tracking points in the tissue region of the subject, and the plurality of ultrasonic images relating to the remaining time phases other than the predetermined time phase, based on the first velocity distribution image Estimating means for estimating a first corresponding point group corresponding to the plurality of tracking points, and estimating a second corresponding point group corresponding to the plurality of tracking points based on the second velocity distribution image, , In each of the time phases, Signal value determining means for determining a signal value in each of the tracking points, the first corresponding point group and the second corresponding point group in the remaining time phases other than the predetermined time phase according to the expansion and contraction of the woven region, and the tracking Generate a first motion information image based on the signal values of the points and the first corresponding point group,
A motion information image generating unit that generates a second motion information image based on the signal values at the tracking point and the second corresponding point group, the first motion information image, and the second motion information image. 1. An ultrasonic diagnostic apparatus comprising: a composite image generating unit that generates a composite image that is combined with a display unit; and a display unit that displays the composite image.

A third aspect of the present invention is to generate a plurality of velocity distribution images for each time phase based on a plurality of ultrasonic images for a plurality of time phases of the subject, and to generate a plurality of velocity distribution images for each time phase. In the ultrasonic image of the predetermined time phase, a plurality of tracking points are set in the tissue region of the subject, and in the plurality of ultrasonic images of the remaining time phase other than the predetermined time phase, the velocity distribution image Based on the estimation of corresponding points corresponding to the plurality of tracking points, in each of the time phases, the signal value at each of the tracking points and the corresponding points is determined according to the expansion and contraction of the tissue region, and the tracking points. And a motion information image is generated based on the signal value at the corresponding point, and the motion information image is displayed.

With such a configuration, an ultrasonic imaging apparatus and a motion information image generating method capable of providing a motion information image representing displacement or strain with high stability even in a living tissue such as a heart. Can be realized.

[0020]

BEST MODE FOR CARRYING OUT THE INVENTION The first and second embodiments of the present invention will be described below with reference to the drawings. In the following description, constituent elements having substantially the same functions and configurations are designated by the same reference numerals, and redundant description will be given only when necessary.

(First Embodiment) In the first embodiment,
Based on the velocity information of the tissue, the noise-reduced motion information is calculated by time integration, and its spatial distribution is provided as an image. The obtained motion information image is displacement image information or distortion image information with respect to a predetermined motion direction.

At this time, in order to effectively obtain a high-quality motion information image for a tissue accompanied by motion, it is important to process velocity information obtained from the tissue while tracking the position of the tissue. . Particularly in the case of a myocardium, the movement amount is large, and tracking of the position is indispensable in order to prevent the target processing position from deviating from the myocardial site (for example, in the heart chamber). Position tracking is also an important factor for obtaining local motion information distribution inside the myocardium.

In order to actually obtain a motion information image (hereinafter referred to as tissue tracking imaging) accompanied by such local tissue position tracking, each processing as shown in FIG. 1 is executed.

[1] Spatio-temporal field of motion (Moti
on-Field: MF) is set (step S
1).

[2] Spatio-temporal distribution image of tissue velocity (Velo)
(city-Field: VF) is obtained (step S
2).

[3] The calculation start time phase is set (step S3).

[4] The region of the tissue (tracking point group) to be traced at the time of the calculation start time is limited (step S4).

[5] The moving position of the tracking point group at each time is V
F · (in the tissue Doppler method, MF is also used) and time intervals are used for successive estimation (step S5).

[6] An input signal is defined at each point of the tracking point group using MF, and integration is performed up to each time (intermediate output point group acquisition; step S6).

[7] At each time, for each point of the output image, the output value is estimated using the values of the nearest intermediate output group around (step S7).

[8] Generation and display of motion information image (step S8).

Here, when a displacement image is obtained, [6]
The displacement can be defined if the input signal is the velocity in the processing of.

When obtaining a distorted image, two types of definitions (Lagrangian str) described below are used.
in, Natural strain)
The ideal distortion can be defined. These definitions temporally track the position of a pair of two points, one point defined in the initial time phase and another point having a predetermined initial length from this point for a one-dimensional rod-shaped tissue model. Obtained by the technology. For more specific details regarding distortion, refer to "Regional Strain and Str".
ain Rate Measurements by
Cardiac Ultrasound: Prince
iples, Implementation and L
imitations "Eur J Echocard
iography (2000) 1,154.70).

[Lagrangian strain]
This is the original definition of strain, and is defined by equation (1) with the initial length L (t0) as a reference. In addition,
FIG. 2A shows the relationship between L (t0) and L (t). _SL (t) = [L (t) -L (t0)] / L (t0) (1) _SL ( t): Lagrangian Strain L (t0): length at reference time t0. Also, Lagrangian strain obtains the displacement by integrating the velocities at each point of the tracked pair, and takes the difference between this pair of displacements. It can be calculated by normalizing with the initial length, and the specific procedure is described below. Figure 2 (b)
Shows the concept.

(1) Two points (points a and b) having an interval of L0
Point) at the t0 time phase. ra (t0), rb (t
0) represents the positions of the points a and b at t0.

(2) L (t) = ra (t) -rb (t) is calculated while temporally tracking the positions of these two points.

(3) SL (t) = (L (t) -L0) /
Calculate L0. Here, L (t) is represented by the following equation (2), and SL (t) is represented by the equation (3-).
It can be obtained by 1) or (3-2).

[0038] L (t) = ∫ Va ( τ) dτ-∫ Vb (τ) dτ + L0 (2) S L (t) = ( track displacement of the tracking displacement -b point of a point) / L0 (3-1) = ∫ [(tracking speed at point a-tracking speed at point b) / L0] dτ (3-2) However, the integration interval is (t0 ≦ τ ≦ t). Also, V
a (τ) and Vb (τ) are tracking velocities at the points a and b.

The integrand of the above equation (3-2) corresponds to the ideal definition of the distortion rate. If this strain rate is output in each time phase without integration, it is also possible to obtain an ideal strain rate after tissue tracking as an output image. However, as described above, since the distortion rate image itself is easily affected by noise, the distortion is focused in this embodiment.

Natural strain as S
_N (t), the length at time t is L (t),
It is formulated as the following equation (4).

S _N (t) = ∫ [(La (x + τ) −L (x)) / L (x)] dx (4) However, the integration interval is (t0 ≦ τ ≦ t).

Further, S _L (t) and S _N (t) are related by the following equation (5).

S _N (t) = ln (1 + S _L (t)) or S _L (t) = exp [S _N (t)]-1 (5) As shown in Expression (4), natural strain
Can be calculated by taking the difference in velocity at each point of the tracked pair, normalizing between the two points to obtain the strain rate, and integrating this strain rate. The specific procedure is described below.
FIG. 3 is a diagram showing the concept.

That is, as shown in FIG. 3, two points (points a and b) having a length of L (t) are set at the time phase t.
Where ra (t) and rb (t) are the time phases t at points a and b.
Represents the position at. Further, it is assumed that the distortion is linear and uniform inside the L (t), and two points (r
1, r2) is set within the length L (t). With such a setting, the strain rate (velocity gradient) between r1 and r2 is N.
Natural strain rate (SRN)
(T)). The SRN (t) is expressed by the following equation (6).

[0045]

[Equation 1]

Using the above (6), S _N (t) is r1
Is calculated by time integration of the velocity gradient (VG (τ)) between t0 and r2 from t0 to t, and S _L (t) = exp (S _N (t)).
S _L (t) can be calculated by the conversion formula of −1. Hereinafter, in the present embodiment, the Lagrangian S
An example using train will be described. As an example related to such tissue tracking imaging, Japanese Patent Laid-Open No.
Although the method of real-time calculation and display of the tissue deformation is described in the publication 001-70303, the specific method of defining the strain is different from this method. That is, in the conventional method, in FIG. 3, distortion based on the instantaneous length is estimated using an assumption,
The ideal distortion is obtained by the expression conversion. Therefore, in order for this result to be equal to the above definition of the ideal distortion, there are the following restrictions due to the fact that the distortion rate between two points having a fixed length in time is obtained. In other words, since only one point is tracked in the conventional method, the assumption that these two points are included in the section having spatially uniform distortion does not hold in the evaluation of the short-axis image of the heart. is there.

Generally, the ultrasonic signal is affected by speckle noise and the like, and in order to obtain a spatially stable distortion rate, this fixed length is increased to some extent (the autocorrelation length of speckle is on the order of several mm). Since it is said that it is 4-1
It is necessary to keep about 0 mm). Therefore, in order to include these two points in the myocardial region that accompanies movement, it can be said that the method of this embodiment in which the two points are directly tracked is advantageous. Even in the conventional example, the accuracy of the distortion rate should be improved by the method of weighted averaging the distortion rate with some changes in the fixed length, and the method of adaptively changing the fixed length spatially (not tracking in time). An attempt is made to do so. However, both are direct 2
The improvement effect is limited due to the fact that there are restrictions such as the above speckles, etc., rather than tracking points.

The processes [4] and [7] are necessary when actually trying to obtain high quality tissue tracking imaging.

That is, [4] is to limit the tissue region. The initial tracking point is set in the tissue, not in the heart chamber. This is because, if the region other than the tissue is not removed from the tracking point group in advance, the value of the non-tissue portion will be added to the output image as noise during the interpolation of the output image. Therefore, it is preferable to limit the tissue area.

[7] is an effective process for estimating a uniform output image even if the tissue is deformed. Figure 4
FIGS. 4A and 4B are diagrams for explaining the processing of [7] using a one-dimensional rod-shaped tissue model for simplicity. As shown in FIG. 4A, when the short portion extends in the movement direction at the calculation start phase, a gap is generated between the positions of the tracking point groups. An appropriate signal value of each point existing in this gap in a predetermined time phase is obtained by weighted addition processing using the value of the nearest intermediate output point surrounding the periphery (that is, the corresponding point in the predetermined time phase described later). To do so. It is preferable that the weighting factor is determined to be larger as the distance between the point of the output image and the positions of the plurality of target tracking point groups is shorter and smaller as the distance is longer. Since such processing can be said to be a kind of interpolation processing, it will be simply referred to as interpolation here. In the example of FIG. 4A, for example, at t2, the signal values at two points existing between the points 1 and 2 are interpolated from the intermediate output values at the points 1 and 2. On the other hand, FIG.
As shown in (b), when a long part shrinks in the calculation start phase, the positions of the point groups to be tracked overlap each other. An appropriate signal value of a point corresponding to this overlap in a predetermined time phase is obtained by the above-described weighted addition process using the value of the intermediate output point (that is, the corresponding point in the predetermined time phase described later) that overlaps. To By such interpolation processing, it is possible to estimate a uniform output image even if the tissue is deformed.

Next, the specific structure and function of the apparatus according to the first embodiment will be described in detail. In the first embodiment, a case will be described in which the heart is targeted and local movement of tissue of the myocardium is evaluated using a two-dimensional image.

FIG. 5 is a block diagram of the ultrasonic diagnostic apparatus according to the first embodiment. The ultrasonic probe 1 includes an ultrasonic transducer array in which a plurality of ultrasonic transducers that convert electric signals into ultrasonic waves are arranged, and the ultrasonic transducer array transmits and receives ultrasonic waves to and from a subject. . In the first embodiment, the ultrasonic probe 1 is a sector probe for the heart.

The transmission unit 2 is for generating a drive signal for transmitting ultrasonic waves from the ultrasonic transducer array, and for each transducer so that an ultrasonic beam is formed toward a predetermined scan line. A drive signal having a predetermined delay characteristic is generated. The receiving unit 3 performs a delay addition process on the ultrasonic echo signals received by each ultrasonic transducer of the ultrasonic transducer array to generate an ultrasonic echo signal corresponding to a predetermined scan line.

The B-mode processing unit 4 performs envelope detection processing on the ultrasonic echo signal subjected to the delay addition processing to generate a B-mode signal corresponding to the amplitude intensity of the ultrasonic echo. The B-mode processing unit 4 also generates a B-mode ultrasonic image that represents a two-dimensional distribution of the B-mode signal on a predetermined cross section. The tissue Doppler processing unit 5 performs quadrature detection processing, autocorrelation processing, and the like, and based on the Doppler shift component of the ultrasonic echo signal subjected to the delay addition processing, the velocity and dispersion of the tissue moving in the subject, Obtain the tissue Doppler signal corresponding to the power. The tissue Doppler processing unit 5 also generates a tissue Doppler ultrasonic image representing a two-dimensional distribution of the velocity, dispersion, and power values on a predetermined cross section.

The motion information processing unit 6 outputs the B output from the B mode processing unit 4 and the tissue Doppler processing unit 5.
Each process for acquiring the motion information image shown in FIG. 1 is executed based on the mode ultrasonic image and the Doppler ultrasonic image. The specific operation of the motion information processing unit 6 will be described in detail later.

The display control unit 7 generates a display image based on the B-mode ultrasonic image, the Doppler ultrasonic image, and the two-dimensional distribution image of displacement or distortion. Examples of the display image include a superimposed image of a B-mode ultrasonic image and a tissue Doppler ultrasonic image, a superimposed image of a B-mode ultrasonic image and a two-dimensional distribution image of displacement or distortion, and the like. The display device 8 displays the display image generated by the display control unit.

The memory 9 stores an ultrasonic image corresponding to each time phase, a velocity distribution image corresponding to each time phase generated by the motion information processing unit 6, and the like.

Next, the specific contents of each processing shown in FIG. 1 of the motion information processing unit 6 will be described in detail with reference to the drawings.

[Generation of Spatio-Temporal Distribution Image of Corrected Velocity Based on Motion Field: Step S and Step S2] First, the motion information processing unit 6 determines the spatio-temporal distribution image of the velocity in the motion direction defined in the motion field 2D distribution image for each phase) is obtained. FIGS. 6A and 6B are explanatory diagrams for defining the playground. FIG. 6A shows a raster motion field (Raster motion field), and FIG. 6B shows a contraction motion field (C).
Each of them represents a playground called an attraction motion field). In each figure, it is the point where the X movement is directed and the direction where the arrow movement is directed.

The raster motion field of FIG. 6A is a motion field along each ultrasonic beam direction of the ultrasonic scan, and the sign of the velocity at this time is positive when it approaches the ultrasonic probe 1, Those that are apart are negative. In such a raster motion field, the observation speed estimated by the tissue Doppler method is, as is well known,
Let θ be the angle between the actual motion direction and the velocity component in the ultrasonic beam direction (called the Doppler angle), and the following equation (7)
It is represented by.

(Observation speed) = cos θ * (actual movement speed) (0 ° ≦ θ ≦ 180 °) (7) The movement dynamics of the tissue in the heart are very complicated, and all are assumed in the raster motion field. Has a problem. In particular, when evaluating a short-axis image that is most suitable for evaluating contraction and expansion in the thickness direction of the myocardium, the 12 o'clock direction (directly above) of the watch when viewed from the contraction center where the ultrasonic beam direction matches the movement direction. ) The movement direction matches only in the extremely narrow range. In the 6 o'clock direction (directly downward direction) when viewed from the center of contraction, the ultrasonic beam direction and the motion direction are parallel, but the directions are opposite.

Therefore, in order to evaluate the motion in the contraction and expansion direction of the myocardium on the short axis, a motion field called a contraction motion field as shown in FIG. 6B is defined, and the direction of the motion field is the actual motion direction. It is preferable to determine the correction speed assuming that there is. In the contraction motion field, the virtual contraction center of the heart wall is set in the image,
The direction towards the virtual center of contraction is defined as the motion field. The setting of the virtual contraction center may be set manually by the operator looking at the displayed ultrasonic image, or may be automatically set by the apparatus from the ultrasonic image by image analysis processing. Is also good.

In the contraction motion field, the velocity component toward the virtual contraction center is obtained by the following equation (8). The sign of the velocity is positive when it approaches the virtual contraction center, and negative when it deviates from it. A two-dimensional distribution image of the correction velocity obtained corresponding to such a contraction motion field is obtained. The two-dimensional distribution image of the corrected tissue velocity may be displayed on the display device 8 if necessary.

(Correction speed) = (Observation speed) / cos θ (0 ° ≦ θ ≦ 180 °) (8) Here, the display of the limit region of the tissue tracking imaging process will be described. In the case of the tissue Doppler method, such a definition of the playground is indispensable for estimating the position of the tracking point cloud described below. This is to determine the moving direction of the tracking point group moving in the two-dimensional image. However, when the Doppler angle θ is close to 90 °, a highly reliable correction speed cannot be obtained.
In the 0 o'clock direction area), there is a limit to the evaluation of motion. Such a region with low reliability is preferably displayed as a limit region of the tissue tracking imaging process so that the user can recognize it. Further, in this limit area, it is desirable not to use the speed information that is the input of the processing. Therefore, a predetermined Doppler angle limit value is set to determine the limit region. The limit value of the Doppler angle (especially in the research stage of the optimum value for practical use) may vary slightly depending on the image quality difference of the tissue Doppler caused by the difference in subject, so that the optimum value The value may be settable by the user. As the initial setting of the Doppler angle limit value, for example, 80 ° to 110 ° is desirable.

It should be noted here that the display of the limit region has another action. This is a point that can suggest what kind of playground is currently set in the image. For example, in the same short-axis image, not only the direction of contraction / expansion (this is the main direction, which is the focus in most cases) but also the motion in the rotational direction (direction parallel to the contour of the myocardium) is the target of evaluation. Therefore, when a plurality of playgrounds can be set, a third person can look at the output image at a glance and recognize what kind of playground. In the case of a rotating motion field, unlike the contraction motion field, a limited region of the Doppler angle occurs near the 12 o'clock direction and 6 o'clock direction of the clock on the short axis.

7 (a) and 7 (b) are display examples of the playground and the limit area. Specifically, FIG. 7A shows a limit area in the contraction motion field, and FIG.
(B) has shown the limit area in a rotary motion field.

In FIGS. 7 (a) and 7 (b), an X mark indicating the virtual center of contraction, an arrow mark indicating the movement direction of the defined motion field, and a limit region (hatched portion) are displayed. For the limit area, a part of the limit area can be grasped by a method of displaying a line at the boundary portion of the limit area, a method of making the color or the brightness in the limit area different from other parts, and the like. These images are displayed on the display device 8 while being superimposed on the above-described two-dimensional distribution image of the corrected tissue velocity.

[Setting of tissue tracking imaging process start phase: Step S3] For a series of tissue Doppler moving images,
From the images stored in the memory, the user selects the time phase t0 (also referred to as the initial time phase t0) at which the processing is started with reference to the electrocardiogram. In this case, not only the start phase of the process, the process at the end of phase (t _ end) is also preferred to leave the selectable by the user.
As a result, the section of the time phase of interest in the evaluation of the cardiac cycle can be set to a desired section (see FIG. 8A).

Further, as shown in FIG. 8B, a predetermined time phase (tR + t_delay time t, where the time phase of the R wave is tR, which is triggered by the R wave of the electrocardiogram at the time of moving image acquisition, t.
_Delay is controllable, and when t_delay = 0 is called R wave synchronization), it may be automatically set as t0 (that is, t0 = tR + t_delay). In this case, it is preferable that the end time phase of the process is exactly one cardiac cycle interval as the timing at which the next trigger occurs, but a predetermined time interval (t_int) from the start time phase is set in advance. If you do
With the time phase of t0 + t_int as the end time phase, it is possible to set the section of the time phase of interest for evaluation of the cardiac cycle to a desired section in the same manner as described above.

[Setting of Calculation Start Time Phase and Setting of Tissue Region (Tracking Point Group) to be Tracked at Calculation Start Time Phase: Steps S3, S4] Next, referring to FIG. This will be described in detail with reference to a). Figure 9
(A) is a flowchart showing a flow of tracking point group setting processing. As shown in FIG. 9A, first, the two-dimensional distribution data IM1 of the corrected tissue velocity in the image area is obtained (step S31). The tracking process may be started for all the regions of the two-dimensional distribution data as it is, but in the case of the heart, the region other than the tissue can also be the tracking target in this case. Therefore, using the brightness information of the B-mode signal, an image IM2 is generated in which only the area exceeding the controllable threshold value is limited to the tissue area (step S3).
2). As a signal source used for such threshold determination, not only the B-mode signal but also power information that can be defined simultaneously with velocity information in the process of tissue Doppler processing may be used. The B-mode signal and the power signal of the tissue Doppler are information that correlates with the signal intensity, and generally, the tissue region has a signal intensity of several 1 as compared with the blood flow region in the heart chamber.
Since it is 0 times larger, it is an effective means for excluding non-tissue regions in the heart chamber from being tracked. When the threshold value is set, the region selected as the effective region (the filled region in FIG. 10) is displayed as the B-mode signal, so that the user can confirm the selected tissue region. It is possible to set an optimum threshold value.

In addition to the above-described region limitation, when setting a region of interest (ROI), this RO
Only the region that exceeds the controllable threshold value for I is limitedly extracted as the tissue region, and the image IM3 or the image IM4 is generated (step S34). In image generation by such ROI setting, the user can perform R
It can be used for the purpose of setting and extracting the OI. In FIG. 10, IM3 uses the circular ROI1 to select the inside of the short-axis epicardium, and IM4 uses ROI2 to select between the short-axis epicardium and the endocardium. The case was shown. If such an ROI is used, for example, the application for removing the papillary muscle region as shown in the figure and the application for removing the intracardiac region when not limited by the B-mode brightness are also applied. Further, since these ROI settings are set for the purpose of extracting the myocardial tissue region, when the above-mentioned setting of the exercise field is set using the ROI of the contour of the heart (MVG-S / L, RO of the method used)
You may use together with I.

[Estimation of the moving position of the tracking point group at each time:
Step S5] Next, for each point of the set tracking point group, the movement position of a series of moving image spaces (that is, a space formed by a plurality of ultrasonic images arranged in time series) Make an estimate.

Here, it is assumed that the moving image is obtained at a constant time interval dt. Also, a contraction motion field based on a virtual contraction center is set as the playing field. For the sake of simplicity, this virtual contraction center is fixed temporally and the motion field is temporally constant.

The method of estimating and tracking the position is preferably the method introduced in TDT. In TDT, accurate tracking is possible when the one-dimensional spatial direction of M mode (that is, the beam direction) is equal to the direction of motion, but in the case of two-dimensional images, the direction of the motion field and the direction in this direction are as follows. Accurate tracking is obtained by using the velocity component.

First, regarding the tracking point group, in the first time phase, the velocity component (correction velocity) in the movement direction of a certain point at that position is Vc, and the position (x,
y) is estimated by the following equation (9).

[0076] (X, y) = (Vc * dt * cos (th), Vc * dt * sin (th)) (9) However, th is an angle formed by the movement direction and the x-axis.

In the next time phase, Vc, th at the position (x, y) is obtained, and the position in the next time phase is estimated using the same equation. By repeating this step, the moving position of this point in each time phase can be obtained.

Note that, in addition to the method using only the information on the current time phase as described above, such a method for estimating the moving position in the future may include information on the current time phase and information on the time phase in the next future. (Eural method: The position in the future obtained above is used as the temporary estimated position 1 and the temporary estimated position 2 in the next future is obtained by the same process at the temporary estimated position 1 and There are also some known methods such as a method of estimating a future position by averaging the estimated position 1 and the temporary estimated position 2. Although the details are omitted, any method is defined by using the velocity information Vc, the movement direction information, and dt, and any estimation method may be used as long as it is in this category.

At this time, it should be noted that the points defined in the image space have a finite size. That is, in the first time phase, the position of the point can be defined on the grit in which the point is definitely defined, but in the subsequent time phases, the estimated position may not always match the grit. In such a case, in order to obtain the value of Vc, it is preferable to interpolate and obtain from the value of Vc defined on the grids of the nearest multiple points around the estimated position.

[Definition of Input Signal of Motion Direction Component and Acquisition of Intermediate Output Point Group by Time Integration: Step S6] Next, on the position of the tracking point group estimated above, the intermediate output for the tissue motion information is output. Make a definition. When the displacement is obtained as the motion information, as described at the beginning, the position of the tracking point group is traced in the time phase direction, and the motion component of the motion direction is integrated by the following equation (10) to calculate the motion. The intermediate output of the displacement with respect to the direction component is obtained on the position of the tracking point cloud in each time phase.

Displacement (x, y, τ) = ΣVc (x, y, τ) * dt (10) where t0 is the first time phase and t is the current time phase. Further, the sum is assumed to be up to t0 ≦ τ ≦ t.

Further, in the case of obtaining the distortion, similarly to the first description, the movement of the velocity is calculated by the following equation (11) while tracing the position of the tracking point group forming a pair of two points in the time phase direction. By integrating the directional components, the ideal intermediate output of the distortion with respect to the directional component is obtained on the position of the tracking point cloud at each time phase. Here, (xa, ya) is one of the tracking point groups selected in the first time phase, and (xb, yb) is a predetermined initial length L0 from the (xa, ya) movement in the first time phase. It is one of the other tracking point groups that form a pair at positions separated in the direction.

[0083]   Distortion (xa, ya, t) = [ΣVc (xa, ya, τ) * dt−ΣVc (xa, ya, τ) * dt] / L0     (10) However, the sum shall be up to t0 ≦ τ ≦ t.

Also when obtaining the values of Vc used for the calculation of these displacements and strains, as in the case of the above-mentioned position estimation, from the values of Vc defined on the grids of the nearest plural points centering on the estimated position. It is preferably obtained by interpolation.

[Estimation of Output Value by Interpolation and Generation and Display of Distorted Image: Steps S7 and S8] First, an image area defining the output image in the current operation time phase is defined, and the output value at each point in the image is It is obtained in steps (1) to (5). Below, the time phase is constant, so the time section is omitted.

(1) At an output point having a predetermined time phase shown in FIG. 11A, as shown in FIG. 11B, a search area having a predetermined shape and size centered on the output point is provided. The shape is preferably round or square, and the size is 5
It is about mm.

(2) As shown in FIG. 11C, the search area is divided into four quadrants, and the tracking point group pi (i = 1 to 4) closest to the output point is searched in each quadrant.

(3) If no trace point group is found in any one quadrant, the output is set to zero.

(4) Nearest tracking point group pi in all quadrants
Is found, the distance Ri between the pi and the output point Ri
And the intermediate output value M (pi) on pi, the output is defined by the following equation (12).

[0090]

[Equation 2]

(5) Perform this step at all output points to obtain output images.

In the first embodiment, each pi
Although the weighting coefficient in 1 is set to 1 / Ri, it may be defined using another function. For example, a function having a larger value as Ri is smaller, such as (1 / Ri) ^1/2 and Gauss function: exp (−σRi ² ), is preferable.

As described above, the displacement or strain image output in a series of steps can be color-converted in the same manner as in the velocity display by the tissue Doppler image, and can be superimposed and displayed on the B-mode image. However, it is preferable to use a color map different from the velocity for such an image of the motion information of the tissue. Specifically, the following color maps are conceivable.

That is, the extending myocardial portion is displayed in red, while the contracting myocardial portion is displayed in blue. In addition, the brightness of red or blue is increased as the magnitude of the strain increases (that is, when the strain has a positive value, red is displayed with the luminance according to the magnitude). On the other hand, when the distortion is a negative value, it is displayed in blue with the brightness according to the magnitude.) At this time, a color bar showing the relationship between the color and the brightness and expansion and contraction is displayed together with the image. It is preferable.

In the color map display, it is not always necessary to perform both expansion and contraction (that is, both red and blue) as described above, and one of them can be selectively displayed as required. It may be. It is preferable that the operator can select which form is used.
When either one is selectively displayed, for example, if only the myocardial stretch is displayed, only the stretched part of the myocardium is displayed in red with brightness according to the magnitude of the strain, and only the contraction of the myocardium is displayed. For example, only the contracted part of the myocardium is displayed in blue with the brightness according to the magnitude of the strain.

Regardless of whether both expansion and contraction are displayed or only one of them is displayed, the expansion and contraction must be distinguished and the time phase must be correctly reflected. An ECG (: electrocardiogram) can be used as a standard for that purpose.

Consider, for example, a case where, in the color map display, one RR cycle of the short axis image of the heart is displayed as a strain image. In this case, for example, during systole, for example, only the extension of the myocardium is displayed, and the thickening of the myocardium is displayed in red. At this time, the signal indicating shrinkage may be determined to be noise and may be erased. Further, for example, in diastole, for example, only the contraction of the myocardium is displayed, and the thinning of the myocardium is displayed in blue. At this time, similarly, the signal indicating expansion may be determined as noise and may be erased.

As described above, the operation time phase of the train and the EC
By imaging the expansion and contraction of the myocardium based on the cardiac time phase by G, it is possible to provide information that is easy for an observer to see and useful for diagnosis and the like.

The color map display described above is realized by the display control unit 7, for example. In addition to the strain image, the color map display is also effective for a displacement image, for example. Also,
From the viewpoint of providing a user-friendly device, it is preferable that the color map display setting can be switched from a short-axis image (for thickening) to a long-axis image (for shorting) by a simple manual operation.

Next, a modification of the first embodiment will be described. In the above description, the virtual center of contraction is fixed for simplicity, but it is also possible to move it within the time phase so as to match the movement of the position of the center of contraction accompanying the movement of the heart. As an example of one means, there is a method in which the virtual contraction center position is set by the user at some specific time phases within the cardiac phase, and the virtual contraction center position at other time phases is obtained by interpolation between the time phases. can give. It is preferable that the interpolation function be linearly estimated from the positions set temporally before and after, or that a function simulating a predetermined motion be provided in advance. As the latter, for example, FIGS. 12 (a), 12 (b), 12 (c)
As shown in, the end diastolic time phase (that is, the time phase ph1)
And two characteristic time phases of the end systole (that is, the time phase ph2) are set, and the virtual contraction center position in these two time phases is set. All of these can be practically effective means for obtaining the virtual contraction center position that matches the movement of the heart while saving the user's setting work as much as possible. This is because if the virtual contraction center setting method is a method in which the user directly places it at a desired position, it can be realized by only placing one point per time phase.

On the other hand, in the conventional MVG-S / L, the direction of motion is defined by the contour. In the case of local disease such as myocardial infarction, there is a case where the movement of contraction and expansion toward the virtual contraction center is not always performed, and this is an attempt to define this in a direction perpendicular to the contour. However, MVG-S / L
Since it is necessary to set the contour, it takes a lot of time and effort for the user to set the contour setting in such multiple time phases. Most ideally, this should be performed by automatic contour tracing, and many attempts have been made, but an error factor depending on ultrasonic image quality is currently a problem.

When evaluating a four-chamber image and a two-chamber image from the apex of the heart, a method of setting a virtual contraction center as shown in FIG. 13 can be considered as a modest method of these. For example, IEICE Transactions D-II Vol. J83-D-
II No. 1 pp. 183-190 (Jan. 2
000)), "Ultrasonic heart wall active contour extraction method using partial shape constrained contour model" is disclosed. According to this method, it is reported that the characteristic structure of the valve annulus can be used for automatic tracking by means of pattern matching. Therefore, by designating two points of the valve annulus and one point of the apex that hardly moves in the image, a predetermined point on the LV-center line in FIG. 13 is set as a dynamic virtual contraction center. Settings are provided.

Further, the Circulation 61:
In 966-972, 1980, "Eval
uation of methods for qua
nitting left ventricula
r segmental wall motion i
n man using myocardial ma
rkers as a standard. According to the document of "", a position 31% from the apex of the heart is preferable as the initial setting of the virtual contraction center.

The processing when a dynamic contraction motion field is set by these methods will be described below. The related process is the step of [estimation of corresponding points in each time phase of the tracking point group] in step S5, and [the definition of the input signal of the movement direction component and the acquisition of the intermediate output point group by time integration] in step S6. Is.

However, even in the case of a dynamic motion field, there is no great change in the basic steps in the case of the static motion field described above. About this situation,
This will be described using 15. The estimation of the tracking position at the time phase t _{i + 1} requires Vc and th at the position (x, y) at the time phase t _i , but these values depend on the motion direction defined at a certain time phase. It can be obtained, and there is no problem even if the direction of movement changes for each time phase. For example, in FIG.
, The estimated position of p (t2) is determined by the motion direction at t1 and Vc corrected in that direction is p (t1).
By applying in, it moves up in the movement direction at t1 (see FIG. 15).
In step S132), at the next t2, th determined by the movement direction at t2 and Vc corrected in that direction are p (t2).
Then, p (t3) is estimated in the movement direction at t2 (step S133 in FIG. 15), and such steps are repeated until the time phase at which tracking is required (step S134 in FIG. 15). Here, it should be noted that the processing procedure itself is the same regardless of whether the movement direction changes between t1 and t2.

Of course, the result of the calculation can be changed between t1 and t2, with or without changing the movement direction. This is a difference in error in the sense that which is more in agreement with the true movement of the tissue. Becomes When the true movement direction changes temporally, it is expected that the error with respect to the tracking position will be improved by changing the movement direction so as to match the change.
Similarly, also for the input signal when obtaining the intermediate output, using Vc corrected in the direction of motion determined by the time phase t is the same as the above-mentioned equation definition.

(Second Embodiment) In the second embodiment,
Tissue tracking imaging is performed by pattern matching of a two-dimensional image. FIG. 16 is a configuration diagram of the ultrasonic diagnostic apparatus according to the second embodiment. The first embodiment is the movement vector processing unit 13 and the motion information processing unit 6
The configurations are different. The other configurations are similar to those of the first embodiment, and thus the description thereof is omitted.

The movement vector processing unit 13 detects the movement position of the tissue by using pattern matching processing between two ultrasonic images having different time phases, and obtains the tissue velocity based on the movement position. Specifically, the partial image in the first ultrasonic image is taken out, and the position of the portion having the highest similarity with the previous partial image in the second ultrasonic image is obtained. By obtaining the distance between the position in the second ultrasonic image and the position of the partial image in the first ultrasonic image, and dividing this distance by the time difference between the first ultrasonic image and the second ultrasonic image, The moving speed of the tissue can be obtained. By performing this process on each point of the ultrasonic image, two-dimensional distribution data of the tissue moving speed can be obtained.

The motion information processing unit 6 obtains a two-dimensional distribution image of displacement or strain of a predetermined section based on the two-dimensional distribution data of the tissue velocity output by the movement vector processing unit 13.

Next, in the ultrasonic diagnostic apparatus according to the second embodiment, each processing shown in FIG. 17 is executed in order to actually obtain the motion information image.

[1] Spatiotemporal distribution image of tissue velocity (Velo
(city-Field: VF) is obtained (step S14).
1).

[2] The calculation start time phase is set (step S142).

[3] Limit the region of tissue (tracking point group) to be tracked in the calculation start time phase (step S14).
3).

[4] The moving position of the tracking point group at each time is V
F · (in the tissue Doppler method, MF is also used) · It is sequentially estimated using time intervals (step S144).

[5] An input signal is defined at each point of the tracking point group using MF, and integration is performed up to each time (intermediate output point group acquisition; step S145).

[6] At each time, the output value is estimated for each point of the output image by using the values of the nearest intermediate output group in the periphery (step S146).

[7] Generation and display of motion information image (step S147).

The contents of each process will be described below.

[Acquisition of Spatiotemporal Distribution Image of Tissue Velocity: Step S141] According to Japanese Patent Laid-Open No. 8-164139.
A two-dimensional spatiotemporal distribution image of tissue velocities is obtained by the movement vector processing unit 13 having processing such as the pattern matching method described in the cross-correlation coefficient display. In pattern matching, the movement position per time interval dt from the current time phase to the next future time phase can be estimated at each point, so if this is divided by dt, a two-dimensional tissue velocity is defined at each point. It

The second embodiment has an advantage over the first embodiment that there is no limit of the Doppler angle and the two-dimensional movement vector is directly estimated, so that there is no assumption of the motion field (operator). This has the advantage that the tracking position can be obtained (even without setting the playing field by, etc.). On the other hand, in order to improve the spatial resolution of velocity estimation, there is a drawback that the calculation time becomes long.

Here, the setting of the motion field is not necessary for the estimation of the tracking position, but in this case as well, in order to shorten the calculation time, as shown in FIGS. 18 (a) and 18 (b), a predetermined value is set. It is preferable to use the definition of the motion field in order to obtain only the desired motion information component for the motion direction.
This point is the same concept as the definition of the input signal of the motion direction component of the first embodiment.

In this case, the purpose of obtaining only the desired motion information component is to make it easier to understand by reducing the dimension of the information amount and providing an output value specialized only for the motion direction component of interest as an image. This is because it is expected that results will be obtained. If two-dimensional tissue velocities are used, two-dimensional displacements and strains can be defined. For example, even if this information is color-converted and superimposed on a two-dimensional B-mode image, it is difficult to understand in practice. Will end up. Rather, for example, it is possible to clearly understand the information of the intended motion direction component by extracting only the component in the direction of the virtual contraction center and displaying it in the same color display as in the first embodiment. Then, when it is desired to evaluate a component in another motion direction, another motion field (for example, a direction of motion orthogonal to the contraction motion field, which is a rotary motion field on the short axis) is set, and an output image is obtained. In this way, the components in the respective movement directions may be separated and evaluated.

Further, since there is no limit on the Doppler angle in the second embodiment, it is not necessary to set or display the limit area in the first embodiment. However, since there is no display of the limit area, in order for a third party to recognize the set playing field, it is necessary to explicitly display the setting state regarding the playing field. In this case, the playing field is indicated by, for example, an icon as shown in FIG. 18 (a) or 18 (b).

[Setting of tissue tracking imaging processing start time phase and setting of tissue region to be tracked (tracking point group) in calculation start time phase: steps S142, S143] These are the same as those in the first embodiment. .

[Estimation of the moving position of the tracking point group at each time:
Step S144] The position of the tracking point group in the next time phase is obtained by multiplying the time phase interval dt by using the two-dimensional tissue velocity in the current time phase.

[Definition of Input Signal of Motion Direction Component and Acquisition of Intermediate Output Point Group by Time Integration: Step S145] The difference between this part and the first embodiment will be described below.

For tracking the position of the tracking point group, the position p (t) of the tracking point group at each time phase t obtained above may be traced, but the tissue velocity at these positions is two-dimensional. Yes, each time phase has its own direction. Therefore, Vc in the first embodiment
The velocity Vc (p of the motion direction component corresponding to (x, y, τ)
In order to obtain (t)), as shown in FIG. 19, the angle formed by the direction of the motion field and the direction of tissue velocity set at the position p (t) of each time phase is bt (p (t)). , And the two-dimensional tissue velocity are V (p (t)) and are calculated by the equation (13).

Vc (p (t)) = | V (p (t)) | * cos [bt (p (T))] (13) First Embodiment Using this angle-corrected velocity component It suffices to perform the same operation as the expression of. In addition, in FIG. 19, in order to have generality, the case where the movement direction dynamically changes was used. However, as described in the first embodiment, it is needless to say that the processing definition is possible even in this case.

[Estimation of Signal Value by Interpolation and Generation / Display of Motion Information Image: Steps S146, 147] The same as in the first embodiment.

By the above procedure, the output image display in the second embodiment can be obtained.

Next, an application example to the time analysis using the output of the tissue tracking imaging according to the present invention will be described.
As shown in the calculation section of the first embodiment, basically output images in a plurality of time phases are provided. Therefore, various time analyzes can be easily applied using the output images in this series of time phases. The main temporal analysis is the time-varying curve of the motion information of the local site and the arbitrary M mode (curved-M).
Also called).

[Application of time-varying curve and ROI tracking technique] ROI according to Japanese Patent Laid-Open No. 10-151133
If an actual example of time analysis using tracking is applied to this embodiment, the time change curve of the motion information of the tissue such as displacement and strain is in a state where the position of the local ROI set on the image matches the position of the myocardium. Provided to enhance diagnostic utility.
It should be noted that, as the ROI position moving means in this case, it is of course possible to apply the actual example shown in the case of moving the setting of the playing field or the actual example of automatic tracking using the velocity information shown in the step of tissue tracking imaging. Absent.

[Curved-M Mode Analysis Application] If the concept of expanded display according to Japanese Patent Laid-Open No. 6-285065 is applied to the present invention, an arbitrary M-mode image of tissue motion information such as displacement and strain can be obtained, and the result of the myocardial throat Since it is possible to temporally compare whether a segment is characteristically in a bad motion state with other segments, it is easy to distinguish at a glance by devising a color map, and diagnostic utility is enhanced. As a color map device, for example, a map in which colors having complementary colors are alternately assigned may be suitable. In this case, it is expected that by setting an appropriate distortion or displacement value in the output range of the map, it will be easier to effectively distinguish between a poorly moving part and a good part.

According to the above-mentioned configuration, even when the target tissue is in motion, by integrating while tracking each position of the tissue to the moving position, the displacement or displacement in the local region of the tissue can be reduced. The distortion is imaged continuously within the sequence of temporal phases of interest. As a result, a distribution image that uses information such as displacement and strain with stability improved by the integration effect is obtained compared to the conventional one, and in the case of the heart, for example, the difference in motion between the endocardial side and the epicardial side of the myocardium You can understand at a glance. Also,
Since these motion information images are continuously obtained within the time phase section, they can be easily applied to the time analysis application.

The present invention has been described above based on the embodiments. However, within the scope of the idea of the present invention, those skilled in the art can come up with various modifications and modifications, and the modifications and modifications. It is understood that the examples also belong to the scope of the present invention. For example, as in the following (1) to (3), various modifications can be made without changing the gist of the invention.

(1) For example, although the case where the received signal is obtained in the two-dimensional space has been described, the same procedure can be applied by extending the dimension even when the received signal is obtained in the three-dimensional space. Is.

(2) The series of processing procedures may be performed separately from the ultrasonic diagnostic apparatus by a general-purpose personal computer or workstation. The series of outputs according to the present invention can be applied not only to the heart but also to the analysis of soft tissues such as the liver and blood vessel walls.

(3) In the generation of the two-dimensional velocity distribution image based on the motion field of the above embodiment, the virtual contraction center is set. However, the virtual contraction center set in this way may deviate from the actual contraction center.

In view of this point, the virtual contraction center is set by the operator or automatically, and points (preferably a plurality of points) near the set virtual contraction center are automatically replaced. It is also possible to have a configuration in which a two-dimensional distribution image having the virtual contraction center of is obtained and combined with these. By using this combined image, it is possible to cancel the influence when the virtual contraction center is displaced, and it is possible to reduce the limit area of the angle correction.

When the other virtual contraction center lies in the limit region, it is preferable that the other virtual contraction center is not adopted and that another virtual contraction center is newly set.

Further, the respective embodiments may be combined as appropriate as much as possible, in which case the combined effects can be obtained. Further, the embodiments include inventions at various stages, and various inventions can be extracted by appropriately combining a plurality of disclosed constituent elements. For example, even if some constituent elements are deleted from all the constituent elements shown in the embodiment, the problem described in the section of the problem to be solved by the invention can be solved, and the effect described in the section of the effect of the invention can be solved. When at least one of the above is obtained, the configuration in which this constituent element is deleted can be extracted as the invention.

[0142]

As described above, according to the present invention, an ultrasonic image device and a motion information image generating method capable of providing a motion information image representing displacement or strain with high stability even in a living tissue such as a heart. Can be realized.

[Brief description of drawings]

FIG. 1 is a flowchart showing a procedure of a motion information image acquisition process executed by an ultrasonic diagnostic apparatus according to a first embodiment.

FIG. 2A is a conceptual diagram for explaining the definition of distortion. FIG. 2B shows Lagrangian s.
It is a conceptual diagram for explaining a train.

FIG. 3 is a conceptual diagram for explaining a natural strain.

4 (a) and 4 (b) are conceptual diagrams for explaining an interpolation process.

FIG. 5 is a schematic configuration diagram of the ultrasonic diagnostic apparatus according to the first embodiment.

6 (a) and 6 (b) are conceptual diagrams for explaining a playground.

7 (a) and 7 (b) are diagrams for explaining a display example of a correction limit region of a Doppler angle.

8A and 8B are conceptual diagrams for explaining the setting of the calculation section.

FIG. 9 is a flowchart showing a flow of tracking point group setting processing.

FIG. 10 is a conceptual diagram for explaining setting of a tracking target tissue region.

11 (a), 11 (b) and 11 (c) are
It is a conceptual diagram for explaining an interpolation process.

12 (a), 12 (b), 12 (c) are
It is a figure for demonstrating the case where a virtual contraction center is temporally moved and set.

FIG. 13 is a diagram for explaining a case of setting a virtual contraction center from anatomical position information of the heart.

FIG. 14 is a conceptual diagram for explaining tracking in a dynamic playground.

FIG. 15 is a flowchart showing a flow of tracking processing in a dynamic playground.

FIG. 16 is a diagram for explaining the schematic configuration of the ultrasonic diagnostic apparatus according to the second embodiment.

FIG. 17 is a flowchart showing a procedure of a motion information image acquisition process executed by the ultrasonic diagnostic apparatus according to the first embodiment.

18 (a) and 18 (b) are diagrams showing display examples regarding setting of a playground.

FIG. 19 is a conceptual diagram for explaining velocity definition in a motion direction component in the second embodiment.

[Explanation of symbols]

1 ... Ultrasonic probe 2 ... Transmission unit 3 ... Receiving unit 4 ... B-mode processing unit 5 ... Organizational Doppler processing unit 6 ... Exercise information processing unit 7 ... Display control unit 8 ... Display device 9 ... Memory 13 ... Moving vector processing unit

Continued front page    F-term (reference) 4C301 CC01 CC04 DD04 DD07 EE11                       FF28 JB28 JC01 JC14 KK02                       KK30 LL03                 5B057 AA09 BA05 CA02 CA08 CA12                       CA16 CB02 CB08 CB12 CB16                       CE08 DA07 DA16 DB02 DB05                       DB09 DC19 DC22 DC36                 5L096 AA03 AA06 BA06 FA14 FA35                       FA67 FA69 HA04 HA05

Claims (30)

[Claims] 1. A storage unit for storing a plurality of ultrasonic images relating to a plurality of time phases of a subject, and a velocity distribution for generating a plurality of velocity distribution images for each of the time phases based on the plurality of ultrasonic images. Image generation means, tracking points setting means for setting a plurality of tracking points in the tissue region of the subject in an ultrasonic image relating to a predetermined time phase among the plurality of ultrasonic images, and residuals other than the predetermined time phase In the plurality of ultrasonic images regarding the time phase of, the estimation means for estimating corresponding points corresponding to the plurality of tracking points based on the velocity distribution image, and in each of the time phases, depending on expansion and contraction of the tissue region. Based on the signal values at the tracking points and the corresponding points, and signal value determining means for determining the signal value at each of the tracking points and the corresponding points,
An ultrasonic diagnostic apparatus comprising: a motion information image generation unit that generates a motion information image; and a display unit that displays the motion information image. 2. The velocity distribution image generation means is defined by the motion field based on the plurality of ultrasonic images based on the plurality of ultrasonic images, and the motion field setting means for setting a motion field for defining the motion direction of the tissue of the subject. The motion velocity distribution image toward the motion direction,
The ultrasonic diagnostic apparatus according to claim 1, further comprising: a distribution image acquisition unit that is obtained for each time phase. 3. The ultrasonic diagnostic apparatus according to claim 2, wherein the motion field is a vector field concentrated at a predetermined one point in the plurality of ultrasonic images. 4. The ultrasonic diagnostic apparatus according to claim 2, wherein the motion field is a predetermined vector field determined from anatomical position information of the tissue in each of the time phases. 5. The ultrasonic diagnostic apparatus according to claim 2, wherein the athletic field is set according to the movement of the region of interest of the tissue in each of the time phases. 6. The ultrasonic diagnostic apparatus according to claim 2, wherein the display device displays, for the sports field, information indicating at least one of a type of a set field and a state of the field. 7. The signal value determining means accumulates speeds at respective time phases of at least a plurality of corresponding points corresponding to the plurality of tracking points based on the speed distribution image for each of the time phases, and performs time integration. Thus, the displacement of the at least the plurality of corresponding points in the latest time phase is determined as a signal value, and the motion information image generating means, based on the displacement of the at least the plurality of corresponding points in the latest time phase, the motion The ultrasonic diagnostic apparatus according to claim 2, wherein a displacement image as an information image is generated. 8. The plurality of tracking points are set as a group of two points separated by an initial length, and the signal value determining means tracks the position of each point constituting each of the two point groups with time. On the other hand, by calculating the motion field direction distance of each of the two point groups in each of the time phases and dividing the motion field direction distance by the predetermined initial length, the strain at the two points separated by the initial length can be obtained. The ultrasonic diagnostic apparatus according to claim 2, wherein a signal value is generated, and the motion information image generation unit generates a strain image as a motion information image based on the signal value related to the strain. 9. The velocity distribution image generation means acquires the movement velocity distribution image by a tissue Doppler method, and the estimation means determines at least the movement velocity toward the movement direction among the movement speeds of the tracking points. The ultrasonic diagnostic apparatus according to claim 2, wherein corresponding points corresponding to the plurality of tracking points in each of the time phases are estimated based on the intervals of the plurality of time phases. 10. The apparatus further comprises Doppler angle setting means for setting an upper limit or a lower limit of the Doppler angle for obtaining a component of the movement velocity of the tissue in the movement direction, wherein the display means shows the set Doppler angle. The ultrasonic diagnostic apparatus according to claim 2, wherein an upper limit or a lower limit is displayed at the same time as the motion information image. 11. The velocity distribution image generating means generates a distribution image concerning the moving velocity of the tissue of the subject by performing pattern matching processing between two ultrasonic images having different time phases, and the estimating means. The ultrasonic diagnostic apparatus according to claim 1, further comprising: estimating corresponding points in each of the remaining time phases based on the respective velocity distribution images and intervals of the plurality of time phases. 12. The signal value determining means is the first of a predetermined time phase.
When the corresponding point of 2 and the corresponding point of 2 overlap at the 3rd corresponding point of the next phase, at least the signal value of the 1st corresponding point and the signal value of the 2nd corresponding point The ultrasonic diagnostic apparatus according to claim 1, wherein the signal value of the third corresponding point is determined based on the above. 13. The signal value determining means, at least when the first tracking point and the second tracking point of the initial time phase overlap each other at a third corresponding point of the next time phase. The ultrasonic diagnostic apparatus according to claim 1, wherein the signal value of the third corresponding point is determined based on the signal value of one tracking point and the signal value of the second tracking point. 14. The signal value determining means determines a signal value of a point in a tissue region near the plurality of corresponding points based on at least the signal values of the nearest corresponding points. 1. The ultrasonic diagnostic apparatus according to 1. 15. The signal value determining means determines the signal values of points in a tissue region near the plurality of corresponding points based on the weighted signal values of the corresponding points. The ultrasonic diagnostic apparatus according to claim 1. 16. The signal value determining means sets the search area of the predetermined size with reference to a point in the tissue area other than the plurality of corresponding points in each of the ultrasonic images in each of the time phases, The corresponding point existing in the search area is searched, and when the corresponding point is searched, the signal value of the point in the tissue area is determined based on at least the signal value of the nearest corresponding point. , If the corresponding points are not searched,
The ultrasonic diagnostic apparatus according to claim 1, wherein the signal value is set to 0. 17. The display means displays the tissue region in a color map according to expansion and contraction.
The ultrasonic diagnostic apparatus described. 18. The apparatus further comprises a function acquisition unit for acquiring a function of time change of tissue motion information in an arbitrary region on the motion information image based on the motion information image, and the display unit acquires the function. 2. The ultrasonic diagnostic apparatus according to claim 1, wherein the function is displayed. 19. The motion information image generating means generates, as the motion information image, an M-mode image relating to predetermined motion information associated with a position on an arbitrary curve or a straight line defined by the user. The ultrasonic diagnostic apparatus according to claim 1. 20. The tracking point setting means sets a threshold value for information correlated with signal intensity at each position of each ultrasonic image, and whether the information correlated with the intensity is larger than the threshold value or not. The ultrasonic diagnostic apparatus according to claim 1, wherein the plurality of tracking points are set based on the determination. 21. The ultrasonic diagnostic apparatus according to claim 1, wherein the display means displays the tracking point set by the tracking point setting unit. 22. The ultrasonic diagnostic apparatus according to claim 1, wherein the tracking point setting means sets the tracking point in a region of interest set in the ultrasonic image of the predetermined time phase. 23. The ultrasonic diagnostic apparatus according to claim 22, further comprising an ROI setting unit for manually setting the ROI. 24. Storage means for storing a plurality of ultrasonic images relating to a plurality of time phases of the subject, and a first contraction center of the tissue of the subject is set for the plurality of ultrasonic images, A contraction center setting means for setting a second contraction center in the vicinity of the first contraction center, and a first motion related to a movement toward the first contraction center based on the plurality of ultrasonic images. Velocity distribution image generation means for generating, for each time phase, a velocity distribution image and a second velocity distribution image relating to the motion along the direction toward the second contraction center; and the plurality of ultrasonic images. Among them, in the ultrasonic image of a predetermined time phase, tracking point setting means for setting a plurality of tracking points in the tissue region of the subject, in the plurality of ultrasonic images of the remaining time phase other than the predetermined time phase , The first velocity distribution image On the basis of,
Estimating a first corresponding point group corresponding to the plurality of tracking points,
Estimating means for estimating a second corresponding point group corresponding to the plurality of tracking points based on the second velocity distribution image; and, in each of the time phases, the tracking points according to expansion and contraction of the tissue region. , The first of the remaining time phases other than the predetermined time phase
Signal value determining means for determining signal values in the corresponding point group and the second corresponding point group, and a first motion information image is generated based on the tracking point and the signal value in the first corresponding point group. A motion information image generating means for generating a second motion information image based on the signal values in the tracking points and the second corresponding point group, the first motion information image and the second motion information image. An ultrasonic diagnostic apparatus, comprising: a composite image generation unit that generates a composite image in which the and are combined, and a display unit that displays the composite image. 25. A plurality of velocity distribution images for each time phase are generated based on a plurality of ultrasonic images of a plurality of time phases of an object, and an ultrasonic wave of the plurality of ultrasonic images for the predetermined time phase is generated. In a sound wave image, a plurality of tracking points are set in the tissue region of the subject, and in the plurality of ultrasonic images related to the remaining time phases other than the predetermined time phase, based on the velocity distribution image, the plurality of tracking points are set. Estimating the corresponding point corresponding to the point, in each of the time phases, determine the signal value at each tracking point and each corresponding point according to the expansion and contraction of the tissue region, the signal value at the tracking point and the corresponding point On the basis of,
Generating a motion information image and displaying the motion information image. 26. In the generation of the velocity distribution image, a motion field that defines a motion direction of the tissue of the subject is set, and based on the plurality of ultrasonic images, the motion direction defined by the motion field is set. Image of the moving velocity distribution toward
The method according to claim 25, further comprising: acquiring for each of the time phases. 27. When generating the velocity distribution image, pattern matching processing is performed between two ultrasonic images having different time phases to generate a distribution image relating to the moving velocity of the tissue of the subject, The motion information image according to claim 25, wherein in estimating points, corresponding points in each of the remaining time phases are estimated based on each of the velocity distribution images and intervals of the plurality of time phases. Generation method. 28. In the determination of the signal value, if the first corresponding point and the second corresponding point of the predetermined time phase overlap at the third corresponding point of the next time phase, then at least the first corresponding point. 26. The ultrasonic diagnostic apparatus according to claim 25, wherein the signal value of the third corresponding point is determined based on the signal value of the corresponding point of 1 and the signal value of the second corresponding point. 29. In the determination of the signal value, the first tracking point and the second tracking point of the initial time phase are the third of the next time phase.
When the corresponding points overlap, the signal value of the third corresponding point is determined based on at least the signal value of the first tracking point and the signal value of the second tracking point. The ultrasonic diagnostic apparatus according to claim 25. 30. In the determination of the signal value, the signal value of a point in a tissue region near the plurality of corresponding points is determined based on at least the signal value of the closest corresponding point. Item 25. The motion information image generation method according to Item 25.