HK1134645B - Method and apparatus for ultrasound synthetic imaging - Google Patents

Description

Method and device for ultrasonic synthetic imaging

Technical Field

The present invention relates to ultrasound synthetic imaging, and in particular to a method and apparatus for medical imaging.

Background

Standard ultrasound imaging

Standard ultrasound imaging includes a sound producing medium for cylindrical waves that are focused at a given point. A complete line of an image can be calculated by dynamic receive beamforming (beamforming) technique using the back-allocated wave of the single sound source. To obtain a complete image, the above process is repeated by sending a set of focused waves that scan along a lateral line at a given depth (called a focal plane). For each focused wave, dynamic beamforming is performed, so that a complete image can be obtained line by line. In the receive mode, dynamic beamforming ensures a uniform focal length. However, in transmit mode, the focus is fixed at a given depth. The final imaging is optimal at the focal plane and over a medium-defined area corresponding to the focal axis length. However, according to the diffraction law, beyond this limited area, at other depths (near and far fields of the focused beam) the imaging quality drops drastically.

To overcome the above limitation, a typical solution is to perform multi-focus imaging: the same image quality is obtained over the entire image range by different transmission focal depths. By each transmission at a given depth of focus, a portion of the image in the imaging region defined by the axial focal length can be obtained. A final image can be obtained by recombining (recombining) the images of the portions corresponding to the different depths. A preferred multi-focus imaging typically requires tens of focal planes. This results in a frame rate limit, typically less than 10 frames/second, which is unacceptable for ultrasound imaging. A good compromise between imaging quality and frame rate is to select about 4 focal depth images.

Ultrasonic synthetic imaging

By synthesizing the dynamic transmission focusing, improvement in the quality of imaging can be expected. The method includes dynamically transmitting focus (i.e., the same number of imaging pixels as the depth of focus) by beamforming re-synthesis, and then combining different sets of sound sources.

Two main embodiments may be referenced: synthetic aperture (Synthetic aperture) and coherent plane wave synthesis.

i) Synthetic pore diameter

In the synthetic aperture method, the ultrasound array is transmitted element by element, and the complete set of impulse responses between each transmit element and each receive element is wave-velocity shaped and recorded, for example, as disclosed in patent document US-668906. Thus, for a composite image that relies on transmitting and receiving each pixel focused in the image, it is possible to obtain by post-processing the data. This document discusses whether the synthesis can achieve better imaging than typical B-mode imaging, and the effects of tissue motion and limited signal-to-noise ratio on the above method. One fundamental problem in synthetic aperture imaging is the low signal-to-noise ratio in the resulting image because a single element is used for transmission. The emission energy of this method is low relative to the full aperture in typical imaging and therefore the penetration depth is limited.

ii) synthetic plane wave technique

Synthetic plane wave imaging methods address, at least in part, the limitations of synthetic aperture imaging. The method comprises transmitting plane waves with different angles in the medium, beamforming the received echo signals, and then combining the different images to resynthesize a final image, as disclosed, for example, in US-6551246. Transmitting plane waves over a complete array can produce a larger pressure field (preassignfield) than synthetic aperture methods. Moreover, the diffraction and attenuation effects of the ultrasound plane wave are much less during propagation in soft tissue than for single element transmission.

The composite dynamic transmission focusing approach approaches the traditional compromise boundary of image quality/frame rate. At higher frame rates (greater than 10 hz), optimal image quality is achieved.

However, the synthetic ultrasound imaging methods currently known, which use the plane wave method, can still be improved in terms of the accuracy of imaging.

Disclosure of Invention

It is an object of the present invention to provide a new method for ultrasound imaging that improves upon current plane wave synthesis ultrasound imaging methods.

To this end, according to one embodiment of the invention, the ultrasound imaging method comprises at least the following steps:

a) a transmission step of transmitting a plurality of ultrasonic waves to an imaging region, a raw data set being respectively obtained by a sensor array respectively responding to each of the ultrasonic waves, the ultrasonic waves having different spatial frequency spectrums for each of a plurality of imaging positions in the imaging region, each of the raw data sets representing time domain signals being obtained by the sensor corresponding to the ultrasonic wave;

b) a coherence enhancing step of synthesizing, for each of a plurality of virtual transmit focal zones in the imaged region, at least one coherent data set from the raw data sets;

c) a beamforming step of obtaining an imaging pixel by beamforming calculation using the coherent data set for each of a plurality of positions included in each of the virtual transmit focal zones.

According to the above steps, prior to beamforming in step b), the spatial coherence of the original data is restored, thus enabling the precise synthesis of the data received from the propagation of the different ultrasonic waves. The necessity to recover the spatial correlation is due to the fact that when illuminating the imaging region with a spatially wide spread wave field, the echoes from the medium can be seen as a wave field generated from incoherent sources (scatterers) randomly distributed over the imaging region: thus, the spatial coherence of the wavefield is lost (or poor) in the raw data.

Then, a more accurate image can be obtained by beamforming the coherent data obtained in the coherence recovery step.

In contrast, in the existing synthetic ultrasound imaging method, the original data is beamformed first and then the images corresponding to the different plane waves are combined, but some information is lost during beamforming and the image combining performed in the prior art cannot restore spatial coherence.

In a variant embodiment of the ultrasound imaging method according to the above-described embodiment of the invention, one and/or other of the following steps may also be included:

-the ultrasonic waves are plane waves with different propagation directions;

-the ultrasonic waves are diffuse waves with different propagation directions;

-the ultrasound wave is a space-time coded excitation;

-said coherence enhancing step operates using a fixed sound speed value;

-said coherence enhancing step comprises an overall sound speed value estimation for said imaged region;

-the coherence enhancing step comprises an overall sound speed value estimation for each imaging position of the imaging region;

-said coherence enhancing step comprises phase difference correction;

-in the coherence enhancing step, each virtual transmit focal area is a line perpendicular to the sensor array;

-said coherence enhancing step comprises:

-a first sub-step of calculating a coherent data set for each of said virtual transmit focal zones by delaying said raw data corresponding to a virtual dynamic transmit focus performed on said virtual transmit focal zone, assuming the same sound speed value in said imaging zone;

-a second sub-step of correcting said delay according to a phase difference estimate of said imaged region, based on said coherent dataset computed in the first sub-step; and by performing virtual dynamic transmit focusing on the virtual focus area, the corrected delay is used to calculate a new coherent data set;

-at least said second sub-step (i.e.: said second sub-step or said first and second sub-steps) is performed several times;

-in said second sub-step, obtaining said phase difference estimate by cross-correlating (cross-correlating) coherent data corresponding to different sensors in each coherent data set.

It is another object of the present invention to provide an apparatus for ultrasound imaging comprising:

-ultrasound transmission means for transmitting a plurality of ultrasound waves to an imaging region, the ultrasound waves having different spatial frequency spectra for each of a plurality of imaging locations in the imaging region;

-raw data obtaining means for obtaining, in response to each of said ultrasonic waves, a respective raw data set;

-combining means for combining at least one coherent RF data set with each raw data set of each of a plurality of virtual transmit focal zones in the imaging region;

-beamforming means for computing a beamforming signal along at least one direction from the coherent data set for each of the plurality of locations in each of said virtual propagation focal zones.

Drawings

Other features and advantages of the present invention will become more apparent upon reading the following detailed description of non-limiting embodiments, with reference to the accompanying drawings.

In the drawings:

FIG. 1 shows a schematic diagram of a synthetic ultrasound imaging apparatus according to one embodiment of the invention;

FIG. 2 shows a block diagram of a portion of the apparatus shown in FIG. 1; and

figure 3 shows a schematic diagram of a composite ultrasound imaging method that can be implemented by the apparatus shown in figures 1 and 2.

Detailed Description

In the drawings, like reference numerals designate identical or similar components.

The apparatus shown in fig. 1 is for synthetic ultrasound imaging of a region 1, for example, for living tissue, in particular, for human tissue of a patient. For example, the apparatus comprises:

an ultrasonic sensor array 2, such as a linear array, typically comprising several tens of sensors (e.g. 100-300) placed side by side along the X-axis, as is known in usual acoustic echo imaging probes (array 2 is suitable for two-dimensional (2D) imaging of said region 1, however array 2 may also be a two-dimensional array suitable for 3D imaging of region 1);

an electronic bay 3 controlling the sensor array and obtaining signals therefrom;

a microcomputer 4 for controlling the electronic bay 3 and displaying the ultrasound images obtained from the electronic bay (in a variant embodiment, a single electronic device may be used to perform all the functions of the electronic bay 3 and the microcomputer 4).

As shown in fig. 2, the electronic bay 3 may include, for example:

n analog/digital converters 5 (A/D)₁-A/D_n) N sensors (T) respectively associated with said sensor array 2₁-T_n) Connecting;

n buffer memories 6 (B)₁-B_n) Connected to the n analog/digital converters 5, respectively;

a central processing unit 8(CPU) communicating with the buffer memory 6 and the microcomputer 4;

a memory 9(MEM) connected to the central processor 8;

a digital signal processor 10(DSP) connected to the central processor 8;

fig. 3 shows an embodiment of the method according to the invention performed using the apparatus shown in fig. 1 and 2, comprising three main steps:

a) transmitting and recording data;

b) synthesis of coherent RF data;

c) and (4) receiving beam forming.

The steps of the method are mainly controlled by a central processing unit 8 and assisted by a digital signal processor 9.

Step a: transmission and data recording

The transducer array is in contact with the imaging medium to be imaged (i.e., the patient's body), and a number N of oblique-plane ultrasound waves are continuously transmitted through transducer array 2 to region 1. The number N of the inclined plane ultrasonic waves may be, for example, 2 to 100. The ultrasonic frequency may be, for example, 0.5 to 100MHz, and more specifically, may be 1 to 10 MHz.

Each tilted plane wave is defined by an angle α, which is the tilt of the propagation direction of the ultrasonic wave with respect to the Z-axis (the Z-axis is perpendicular to the X-axis of the linear transmission array 2 and defines an imaging plane with the X-axis).

Each oblique plane wave encounters some scatterers (speckles) and the echoes are scattered to sensor array instance 2, so that each sensor T in the array₁-T_nAn echo scatter signal is received. Then, the echo scattered signals received by the n sensors are digitized by the analog/digital converters 5, respectively, and stored in the n buffer memories 6. After transmitting one plane wave, the data stored in the n buffer memories is called raw RF data. Thus, the raw RF data may be viewed as a matrix representing the time signals received by all of the sensors in the array 2 after transmission of a plane wave. "RF" is a general term in the art and refers to the frequency of ultrasound waves (typically in the range of 0.5 to 100MHz), although it is not intended to be limiting in any way.

It is noted that the ultrasonic plane waves can also be replaced by divergent ultrasonic waves having different directions of propagation.

In all embodiments, the ultrasound waves may be space-time coded (spatio-temporal coded), for example, multiple ultrasound waves of different directions may be propagated and processed simultaneously.

Step b: synthesis of coherent RF data

M coherent composite RF data matrices can be calculated by using the extended time delay and summation operation through N obtained original RF data matrices. The computed M matrices correspond to the discrete echoes obtained from a given virtual dynamic transmission focal line. It can be obtained by the following procedure.

1) Substep b.1: assuming that the speed of propagation of the acoustic wave through the medium is constant, a first set of coherent RF data is synthesized.

For simplicity, we agree here that the wavefield is transmitted from a sensor at time t-0 and location (x-0, z-0). Consider a virtual propagation focal position F (x)₁Z). If the medium is insonified with a plane wave of inclination α, the wave reaches a position F (x) in the medium₁Z) the time required is:

τ_ec(α，x₁，z)＝(zcosα+x₁sinα)/c (1)

wherein x₁And Z is the distance from the X and Z axes, and c is the speed of sound in region 1. The speed of sound c may be a predetermined value, or an overall estimate.

The time required to return to the position of a given sensor on the X-axis is:

thus, the total travel time τ for a given plane wave transmission is:

for the virtual transmission focus position F (x) considered₁Z) which gives the link between the raw RF data (RFraw (x, z, α) and spatially coherent RF data:

where B (α) is a weighting function for each angular distribution. A virtual focus line (virtual focus line) is defined as being at the same side position x₁All virtual focus points (x) F (x) of (same spatial position)₁Z). Along a given line x₁Each virtual focal line of (a) is represented by a 2D matrix of coherent RF data: matrix RFcoherent (x)₁，x，z)。

M of said coherent RF data matrices are calculated, M being the number of virtual transmission focal lines to be calculated (i.e. x)₁M values of).

For example, M may correspond to the number of sensors in the array 2, or may be a larger number.

2) Substep b.2: by the phase difference correction, a corrected focusing rule of the medium is determined.

In the previous calculations, the potential aberrations of the medium that may cause distortion of the ultrasound propagation are not taken into account. Those local aberrations may consist of local variations in acoustic properties, such as sound velocity, density or sound absorption (in medical ultrasound, such heterogeneity exists in sound velocity, roughly ranging from 1460 m.s. for fat^-1To 1560m.s for muscle^-1). Such phase differences cause errors that degrade the spatial coherence of the composite signal and the quality of the resulting ultrasound imaging.

In order to correct the above-described error, a known offset correction method may be used. By aligning the line x₁The methods are applied to coherently synthesized RF data obtained above using a set of time delays δ (x)₁X, z) can be estimated and added to the different travel times:

τ_rew(α，x₁，x，z)＝τ(α，x₁，x，z)+δ(x₁，x，z) (5)

where δ is the retardation correction for the error introduced by assuming cylindrical focusing laws.

By τ_new(α，x₁X, z) as τ (α, x) in formula (4)₁X, z) is recalculated for the M coherent composite data matrices in step b.1.

These phase deviation corrections are the same as the local estimation of the sound velocity at each imaging position in the region 1.

The phase difference correction method relies on the spatial coherence of the recorded wavefields.

The spatial coherence of one of the wavefields is determined by its spatial covariance (space covariance). It determines the correlation cross-correlation between values at two points in the field as a function of their separation with respect to the two points. The correlation function corresponds to second order statistics of spatial perturbation (space perturbation) of the wavefield produced by an incoherent source.

One of the important theorems of optics, the so-called Van cite-Zernike (Van citter-Zernike) theorem, describes the second order statistics for such sites. The Van Satt-Zernike theorem describes two points X on the observation plane₁And X₂Field spatial covariance of (a), and at spatial frequency (X)₂-X₁) The Fourier transform of the source aperture function is equal for λ z, where λ is the wavelength and z is the distance from the source to the viewing surface.

As a direct result, the sharper the incoherent source point (sharer), the larger the region where the fields sensed at two different points on the viewing surface have a high degree of similarity.

The phase difference correction technique used in medical ultrasound imaging is mainly based on this conclusion. Of course, if the field recorded by the sensor array corresponds to discrete echoes from an incoherent source set located at a small spatial spot, there is a high similarity between the field received at one element in the array and the field received at the adjacent element to that element. Thus, a simple correlation between signals received at adjacent elements allows the time delay between these signals to be recovered. Applying this method to the elements in all arrays, a set of time delays between all elements is obtained, so that the phase difference medium can be completely described (phase difference correction techniques make the assumption that a phase difference layer is thin and close to the array, and only time shifts (known as "phase screen approximation") are introduced in the elements of the array).

3) Finally, the substep b.2 is repeatedly carried out

Finally, sub-step b.2 may be repeatedly performed. The number of repetitions may be predetermined or may be stopped when the error value falls below a predetermined threshold. For example, the error value is: e ═ Σ δ (x)₁，x，z)²Or otherwise.

Instead of just repeating sub-step b.2, it is also possible to repeat sub-steps b.1 and b.2 simultaneously for optimizing the delay rule estimation.

Step c: receive beamforming

After step b) is completed, a receive beamforming is performed for each of the M coherent RF data matrices to compute the final ultrasound imaging. The delay rule used is calculated by a phase difference correction method:

by coherently adding the contributions of each scatterer, one point (x) in the image can be obtained₁Z), i.e. using τ_new(x₁X, z) delay signalAnd adds them in the array direction X:

where A is the receive apodization function (receive apodization function) as the forming line x in the final imaging₁Is a function of x.

Thus, the imaging includes M lines.

Alternative embodiment

By slightly optimizing the foregoing embodiments, the variance of the phase difference distortion estimates may be improved.

In step b, each of the M matrices calculated corresponds to a discrete echo resulting from a given virtual dynamic propagation focal line obtained by summing the raw per-channel data at different angles α.

However, in some embodiments, step b is performed to synthesize a composite corresponding to a given line x₁K independent versions of the same matrix.

For example, such a separate version can be obtained very simply with a different and reduced set of angles α for each version.

The other acquisition corresponds to a given line x₁The method of K independent versions of the same matrix, comprising the modification step a: for successive subsets of the array sensors (i.e., successive sub-apertures), N tilted plane waves are transmitted. Also, in step 2, the set of N raw R F data corresponding to line x is obtained using the N data obtained by the different sub-apertures of the array₁K versions of the virtual matrix of (1)。

Thus, in step b.2, δ (x) can be estimated for the phase difference distortion₁X, z) is improved because it corresponds to a virtual line x₁Of the same matrix, these phase differences must be the same. Thus, by δ (x) for K different versions of the same matrix₁X, z) simple averaging, the estimated variance can be reduced.

More complex recombination techniques may also be used, such as DORT techniques (see, e.g., Prada C, Thomas JL. Experimental subset OF localization OF satellites. side decomposition OF THE time transformed operator expressed as a covarian. JOURNAL OF THE ACOUSTICAL SOCITY OF AMERICA 114 (1): 235. 243 JUL 2003,

and Prada C, Manneville S, spirolansky D, et al. composition of the time reversal operator: detection and selective focusing on two scatter. JOURNAL OF THE ACOUSTICAL SOCIETY OF AMERICA 99 (4): 2067-: year 7 months of 235-2432003;

and Prada C, Manneville S, Spoliansky D, et al, temporal inverse operator decomposition: detection and selection focused on double scatterers, proceedings of the american society of acoustics 99 (4): 2067 year 20761996 month 4).

The method according to the invention can also be used, for example, for:

-for at least one 2D or 3D ultrasound imaging with dynamic focusing on transmission and reception;

-a set of ultrasound images for dynamic focusing on transmit and receive, which sets can be incoherently added to produce a composite image;

-at least one 2D or 3D ultrasound imaging for dynamic focusing on transmission and reception, and one 2D or 3D color flow image;

-for at least one 2D or 3D high quality ultrasound tissue harmonic imaging;

-for at least one 2D or 3D high quality ultrasound contrast imaging using injected contrast agent;

at least one 2D or 3D ultrasound imaging for dynamic focusing on transmit and receive, combined with other typical modes, such as focused color flow imaging, or harmonic imaging.

Claims (13)

1. A method for ultrasound imaging, characterized by at least the following steps:

a) a transmission step of transmitting a plurality of ultrasonic waves to an imaging region, a raw data set being respectively obtained by a sensor array respectively responding to each of the ultrasonic waves, the ultrasonic waves having different spatial frequency spectrums for each of a plurality of imaging positions in the imaging region, each of the raw data sets representing time domain signals being obtained by the sensor corresponding to the ultrasonic wave;

b) a coherence enhancing step of synthesizing, for each of a plurality of virtual transmit focal zones in the imaged region, at least one coherent data set from the raw data sets;

c) a beamforming step of obtaining an imaging pixel by beamforming calculation using the coherent data set for each of a plurality of positions included in each of the virtual transmit focal zones.

2. The method of claim 1, wherein the ultrasonic waves are plane waves having different propagation directions.

3. The method of claim 1, wherein the ultrasonic waves are diffuse waves having different propagation directions.

4. The method of claim 1, wherein the ultrasound wave is a space-time coded excitation.

5. The method of claim 1, wherein the coherence enhancing step operates using a fixed speed of sound value.

6. The method of claim 1, wherein the coherence enhancing step comprises an overall sound speed value estimate for the imaged region.

7. The method for ultrasound imaging according to claim 1, wherein the coherence enhancing step comprises an overall sound speed value estimation for each imaging position of the imaging region.

8. The method for ultrasound imaging according to claim 1, wherein the coherence enhancing step comprises phase difference correction.

9. The method of claim 1, wherein in the coherence enhancing step, each of the virtual transmit focal zones is a line perpendicular to the transducer array.

10. The method for ultrasound imaging according to claim 1, wherein the coherence enhancing step comprises:

-a first sub-step of calculating a coherent data set for each of said virtual transmit focal zones by delaying said raw data corresponding to said virtual dynamic transmit focusing performed on said virtual transmit focal zones, assuming the same sound speed values in said imaging zones;

-a second sub-step of correcting said delay according to a phase difference estimate of said imaged region, based on said coherent dataset computed in the first sub-step; and the corrected delay is used to calculate a new coherent data set by performing the virtual dynamic transmit focusing on the virtual focus area.

11. A method for ultrasound imaging according to claim 10, wherein the second sub-step or the first and second sub-steps are performed several times.

12. A method for ultrasound imaging according to claim 10, wherein in the second sub-step the phase difference estimate is obtained by cross-correlating the coherent data corresponding to different ones of the sensors in each of the sets of coherent data.

13. An apparatus for ultrasound imaging, comprising:

-ultrasound transmission means for transmitting a plurality of ultrasound waves to an imaging region, the ultrasound waves having different spatial frequency spectra for each of a plurality of imaging locations in the imaging region;

-raw RF data obtaining means for obtaining, in response to each of said ultrasonic waves, a respective raw RF data set;

-synthesizing means for synthesizing at least one coherent data set by each raw data set of each of a plurality of virtual transmit focal zones in the imaging region;

-beamforming means for computing a beamforming signal along at least one direction from the coherent data set for each of the plurality of locations in each of said virtual propagation focal zones.