Universal B-ultrasonic imaging spatial resolution modeling method and system

Abstract

The invention relates to the technical field of ultrasonic medical images, and discloses a universal B-ultrasonic imaging spatial resolution modeling method and system. The method comprises the following steps: configuring ultrasonic transducer parameters; calculating a shock wave signal of a single emission array element at any specific scattering point, superposing all shock wave signals to generate a total shock wave signal, and multiplying the total shock wave signal by the reflectivity of the scattering point to obtain an initial echo signal; echo signals reaching the receiving array elements are calculated; performing linear superposition on all the echo signals to obtain a total echo signal; performing image reconstruction on the total echo signal to output a B ultrasonic image; obtaining a spatial resolution quantification result through a point spread function; and changing parameters of the ultrasonic transducer, and calculating a spatial resolution quantification result to realize modeling of an influence mechanism of different parameters on the spatial resolution. The method can support modeling of the ultrasonic transducer with any array arrangement, any array element direction and any emission waveform and any pulse echo sequence, and has a high degree of freedom.

Description

Technical Field

[0001] The present invention relates to the technical field of ultrasonic medical imaging, and particularly relates to a general B-mode ultrasound imaging spatial resolution modeling method and system. Background Art

[0002] B-mode ultrasound imaging is a radiation-free, non-invasive and highly real-time medical imaging technology. It emits ultrasonic waves into a living body by controlling an ultrasonic transducer and collects echo signals to reconstruct the structural information in the living body. Spatial resolution is an important indicator for measuring B-mode ultrasound imaging, which directly affects the efficiency and accuracy of diagnosis. Therefore, improving spatial resolution is one of the key objectives in the design and development of B-mode ultrasound imaging systems.

[0003] For B-mode ultrasound imaging, numerous factors in the imaging chain will affect the spatial resolution, such as array geometry, ultrasonic transducer frequency response, and pulse-echo sequence. Currently, there is a lack of a unified model to study the influence mechanism of these factors on spatial resolution. Therefore, it is of great significance to develop a general spatial resolution research model.

[0004] In the prior art, there is a spatial resolution research model that simulates specific imaging scenarios based on the Field Ⅱ toolkit, such as a research model on the influence mechanism of ultrasonic transducer factors on spatial resolution based on plane-wave imaging. This model uses the Field Ⅱ ultrasonic simulation toolkit to simulate the imaging of the plane-wave sequence of a linear array ultrasonic transducer, and studies the influence of these system parameters on spatial resolution by changing the parameters of the ultrasonic transducer, such as the center frequency, bandwidth, array width, etc. Reference: ALOMARI Z, HARPUT S, HYDER S, et al. The effect of the transducer parameters on spatial resolution in plane-wave imaging[C] / / IEEE International Ultrasonics Symposium (IUS), Taipei, China: IEEE, 2015: 1-4.

[0005] In the prior art, there is also a spatial resolution research model based on an analytical formula. This model establishes an analytical expression of the point spread function of the imaging system through the approximation of the Lippmann-Schwinger equation and studies the spatial resolution based on the point spread function. The influence of these system parameters on the spatial resolution can be studied by changing the number of ultrasonic transducer array elements, the array width, and the bandwidth. Reference: ROQUETTE L, SIMEONI M, HURLEY P, et al. On an analytical, spatially-varying, point-spread-function[C] / / IEEE International Ultrasonics Symposium (IUS), Washington, USA: IEEE, 2017: 1-4.

[0006] The above research model on the influence mechanism of ultrasonic transducer factors on spatial resolution based on plane wave imaging has many limitations. First, the spatial resolution can only be studied based on the plane wave sequence. However, B-mode imaging also has multiple pulse echo sequences, such as focused line scan imaging, coherent plane wave compound imaging, and synthetic aperture imaging, etc. The research on the spatial resolution based on other imaging sequences is also extremely important. Second, it can only study the influence mechanism of ultrasonic transducer parameters on the spatial resolution and cannot study the influence mechanisms of other system parameters, such as sequence parameters, beam synthesis algorithm parameters, etc. Third, it can only perform simulation imaging on simple ultrasonic transducer arrays, such as linear arrays, and complex ultrasonic transducer arrays, such as circular arrays and spherical arrays, cannot be used. Summary of the Invention

[0007] To solve the above technical problems, the present invention provides a general B-mode imaging spatial resolution modeling method and system, which can be applicable to various scenarios and support the research on the influence mechanisms of multiple system parameters on the spatial resolution; it can guide the optimization of system parameters in the B-mode imaging system to improve the spatial resolution of imaging.

[0008] To solve the above technical problems, the present invention adopts the following technical solutions: In a first aspect, the present invention provides a general B-mode imaging spatial resolution modeling method, including: Configuring ultrasonic transducer parameters; Calculating the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point through the transmitting waveform, transmitting apodization coefficient, directivity function, and transmitting delay, and superimposing the shock wave signals of all transmitting elements to generate a total shock wave signal; multiplying the total shock wave signal by the reflectivity of the scattering point to obtain the initial echo signal of the scattering point; ​Calculate the echo signal after the initial echo signal of a single scatterer reaches the receiving array elements based on the received waveform, receive apodization coefficient, directivity function, and receive delay; linearly superimpose the echo signals of the initial echo signals of all scatterers reaching the current receiving array element to obtain the total echo signal of the current receiving array element, and further obtain the total echo signals of all receiving array elements; Perform image reconstruction on the total echo signal and output a B-mode ultrasound image; Set any point within the imaging area as an ideal scatterer, obtain the point spread function by calculating the B-mode ultrasound image of the ideal scatterer, take the profile line of the point spread function along a preset direction, and the full width at half maximum value of the profile line is the quantization result of the spatial resolution in this direction; By changing the ultrasonic transducer parameters and calculating the corresponding spatial resolution quantization results, model the influence mechanism of different parameters on the spatial resolution.

[0009] In one embodiment, the configurable ultrasonic transducer parameters include one or more of the number of array elements in the array, element width, element pitch, element height, element curvature, array width, array arrangement method, and the frequency response of the ultrasonic transducer; the frequency response includes the center frequency and the relative bandwidth, and the relative bandwidth is the ratio of the difference between the upper cut-off frequency and the lower cut-off frequency to the center frequency.

[0010] In one embodiment, the calculation of the shock wave signal of a single transmitting array element of the ultrasonic transducer at an arbitrary specific scatterer based on the transmitted waveform, transmit apodization coefficient, directivity function, and transmit delay specifically includes: The calculation method of the shock wave signal of each transmitting array element at the scatterer is: the transmitted waveform is multiplied by the transmit apodization coefficient of the transmitting array element, multiplied by the directivity value of the directivity function at the scatterer, and then multiplied by the attenuation function of the ultrasonic wave; Among them, the attenuation function of the ultrasonic wave is determined according to the propagation distance from the array element to the scatterer; Determine the position of the shock wave signal in the time domain through the transmit delay.

[0011] In one embodiment, the transmitted waveform is generated according to the frequency response of the ultrasonic transducer, and the transmit apodization coefficient is the weight coefficient of each transmitting array element; the directivity function is determined by the element width and the center frequency.

[0012] In one embodiment, the calculation of the echo signal after the initial echo signal of a single scatterer reaches the receiving array element based on the received waveform, receive apodization coefficient, directivity function, and receive delay specifically includes: The calculation method of the echo signal arriving at the receiving element is as follows: the received waveform applies the ultrasonic transducer frequency response to the initial echo signal of the scattering point, multiplies by the receiving apodization coefficient of the receiving element, multiplies by the directivity value of the directivity function at the scattering point, and then multiplies by the attenuation function of the ultrasonic wave; Among them, the attenuation function of the ultrasonic wave is determined according to the propagation distance from the scattering point to the receiving element; The position of the echo signal in the time domain is determined by the reception delay.

[0013] In one embodiment, performing image reconstruction on the total echo signal and outputting a B-mode ultrasound image specifically includes: Performing time gain compensation, band-pass filtering, Hilbert transform envelope detection, delay focusing beam synthesis, and logarithmic compression on the total echo signal in sequence, and outputting a B-mode ultrasound image.

[0014] In one embodiment, obtaining the point spread function by calculating the B-mode ultrasound image of the ideal scattering point specifically includes: Denote the pixel value of any pixel point in the B-mode ultrasound image corresponding to the ideal scattering point as the function f(x, y), where x and y are the row number and column number of the pixel point respectively, and the function f(x, y) is the point spread function.

[0015] In a second aspect, the present invention provides a general B-mode ultrasound imaging spatial resolution modeling system, including: Parameter configuration module: used to configure ultrasonic transducer parameters; the ultrasonic transducer parameters include one or more of the number of elements in the array, element width, element pitch, element height, element curvature, array width, array arrangement method, and ultrasonic transducer frequency response; Signal emission module: calculates the shock wave signal of a single emission element of the ultrasonic transducer at any specific scattering point through the emission waveform, emission apodization coefficient, directivity function, and emission delay, and superimposes the shock wave signals of all emission elements to generate a total shock wave signal; multiplies the total shock wave signal by the reflectivity of the scattering point to obtain the initial echo signal of the scattering point; Signal reception module: calculates the echo signal after the initial echo signal of a single scattering point arrives at the receiving element through the reception waveform, reception apodization coefficient, directivity function, and reception delay; linearly superimposes the echo signals of the initial echo signals of all scattering points arriving at the current receiving element to obtain the total echo signal of the current receiving element, and further obtains the total echo signals of all receiving elements; Image reconstruction module: performs image reconstruction on the total echo signal and outputs a B-mode ultrasound image; Resolution quantization module: Set any point within the imaging area as an ideal scattering point, obtain the point spread function by calculating the B-mode ultrasound image of the ideal scattering point, take the profile line of the point spread function along a preset direction, and the full width at half maximum value of the profile line is the quantization result of the spatial resolution in this direction; Modeling module: By changing the ultrasonic transducer parameters and calculating the corresponding spatial resolution quantization results, the modeling of the influence mechanism of different parameters on the spatial resolution is realized.

[0016] In one embodiment, calculating the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point through the transmitting waveform, transmitting apodization coefficient, directivity function, and transmitting delay specifically includes: The calculation method of the shock wave signal of each transmitting element at the scattering point is: the transmitting waveform is multiplied by the transmitting apodization coefficient of the transmitting element, multiplied by the directivity value of the directivity function at the scattering point, and then multiplied by the attenuation function of the ultrasonic wave; Wherein, the attenuation function of the ultrasonic wave is determined according to the propagation distance from the element to the scattering point; The position of the shock wave signal in the time domain is determined by the transmitting delay.

[0017] In one embodiment, calculating the echo signal after the initial echo signal of a single scattering point reaches the receiving element through the receiving waveform, receiving apodization coefficient, directivity function, and receiving delay specifically includes: The calculation method of the echo signal reaching the receiving element is: the receiving waveform applies the frequency response of the ultrasonic transducer to the initial echo signal of the scattering point, multiplied by the receiving apodization coefficient of the receiving element, multiplied by the directivity value of the directivity function at the scattering point, and then multiplied by the attenuation function of the ultrasonic wave; Wherein, the attenuation function of the ultrasonic wave is determined according to the propagation distance from the scattering point to the receiving element; The position of the echo signal in the time domain is determined by the receiving delay.

[0018] Compared with the prior art, the beneficial technical effects of the present invention are: For the generation part of the ultrasonic signal, the signal acquisition model proposed by the present invention can support the modeling of ultrasonic transducers with any array arrangement, any element orientation, any transmitting waveform, and any pulse echo sequence, with a high degree of freedom.

[0019] Compared with the spatial resolution research model based on analytical formulas, the research model based on numerical simulation of the present invention has lower complexity and computational amount. Compared with the spatial resolution research model based on the Field Ⅱ toolkit, the present invention has higher flexibility and supports the modeling of any ultrasonic transducer and any pulse echo sequence. Description of the Drawings

[0020] Figure 1 is the method flow chart in the embodiment of the present invention; Figure 2 is the specific flow chart for digital modeling of the B-ultrasound imaging system in the embodiment of the present invention; Figure 3 is the schematic diagram of some parameters that can be customized for the linear array ultrasonic transducer in the embodiment of the present invention; Figure 4 are the frequency domain spectrum and time domain spectrum of the transmitted ultrasonic wave under the set center frequency and bandwidth in the embodiment of the present invention; Figure 5 is the schematic diagram of the principle of the ultrasonic signal generation model based on linear superposition in the embodiment of the present invention; Figure 6 are the modeling results of the point spread function and the quantization results of the spatial resolution of ultrasonic transducers with different geometric shapes in the first embodiment; Figure 7 are the modeling results of the point spread function and the quantization results of the spatial resolution of ultrasonic transducers with different center frequencies in the second embodiment. Specific Embodiments

[0021] A preferred embodiment of the present invention will be described in detail below with reference to the accompanying drawings.

[0022] As Figure 1 shown, the present invention provides a general method for modeling the spatial resolution of B-ultrasound imaging, including the following steps: S1, Configure the parameters of the ultrasonic transducer; S2, Calculate the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point through the transmitted waveform, transmit apodization coefficient, directivity function, and transmit delay, and superimpose the shock wave signals of all transmitting elements to generate a total shock wave signal; Multiply the total shock wave signal by the reflectivity of the scattering point to obtain the initial echo signal of the scattering point; S3, Calculate the echo signal after the initial echo signal of a single scattering point reaches the receiving element through the received waveform, receive apodization coefficient, directivity function, and receive delay; Linearly superimpose the echo signals of the initial echo signals of all scattering points reaching the current receiving element to obtain the total echo signal of the current receiving element, and further obtain the total echo signals of all receiving elements; S4, Perform image reconstruction on the total echo signal and output a B-ultrasound image; S5, Set any point in the imaging area as an ideal scattering point, obtain the point spread function by calculating the B-ultrasound image of the ideal scattering point, take the profile line of the point spread function along the preset direction, and the full width at half maximum value of the profile line is the quantization result of the spatial resolution in this direction; S6. By changing the ultrasonic transducer parameters and calculating the corresponding spatial resolution quantization results, the modeling of the influence mechanism of different parameters on the spatial resolution is realized.

[0023] The present invention supports digital modeling of B-ultrasound imaging systems under various parameter settings. The specific process of modeling is as Figure 2 shown. First, set the ultrasonic waveform to be emitted by the ultrasonic transducer. The ultrasonic transducer performs the transmission and reception of ultrasonic signals according to the set pulse-echo sequence, where the pulse-echo sequence includes a transmission sequence and a reception sequence. Then, encode and arrange the original digital ultrasonic signals according to the execution order of the pulse-echo sequence and store them. Next is signal processing. After obtaining the digital ultrasonic signals, since the sound wave will attenuate as it propagates in space, intensity compensation needs to be performed according to the sound wave propagation time, which is called time gain compensation. Then, perform band-pass filtering on the signals to eliminate the interference of clutter signals outside the ultrasonic transducer bandwidth and system noise. Subsequently, the Hilbert transform needs to be performed on the signals. The original signal is used as the real part, and the transformed signal is used as the imaginary part, which is called envelope detection. After that is image reconstruction. Calculate the time index corresponding to each pixel in each pulse-echo sequence in the signal, which is called delay focusing. According to the calculated delay values, index the corresponding intensity values for the real and imaginary parts of the signal respectively, and then perform weighted summation. This process is called beamforming. Each pixel of the image obtained by beamforming is a complex value. Taking the modulus length of it can obtain the reconstructed pixel value of each pixel. In order to see the structures with relatively weak echo signals inside biological tissues, image compression is required. The commonly used method is logarithmic compression.

[0024] Logarithmic compression first normalizes the image pixel values, then takes the common logarithm of each pixel value, and then multiplies by 20. The calculation formula is as follows: ; In the formula, is the original pixel value of the i-th pixel of the image, is the maximum original pixel value, is the pixel value after compressing the i-th pixel of the image, with the unit of dB. The range of the compressed pixel values is (-∞, 0]. Therefore, a lower limit threshold needs to be selected, and the pixel values less than this threshold are all set equal to this threshold. For example, if -60 is selected as the lower limit threshold, the dynamic range of the image is said to be 60 dB. By selecting different lower limit thresholds, different degrees of image compression can be achieved. Finally, after obtaining the B-ultrasound image, image display is performed. By simulating imaging of ideal scatter points to obtain the point spread function image of the imaging system, the quantization of the spatial resolution can be realized.

[0025] The specific process of the present invention is divided into three steps: ultrasonic signal acquisition, ultrasonic image reconstruction, and spatial resolution calculation.

[0026] The first step is the acquisition of ultrasonic signals, also known as the forward process, which refers to the process of controlling the ultrasonic transducer to emit ultrasonic waves and collect echoes. Based on a linear model, digital modeling of point target echo signals is realized. The core process includes three key links: ultrasonic transducer definition, ultrasonic signal generation, and ultrasonic signal storage.

[0027] Link 1, ultrasonic transducer definition: The present invention supports customizing various parameters of the ultrasonic transducer. Taking the common linear array ultrasonic transducer as an example, the parameters that can be customized are as Figure 3 shown. It includes: the number of array elements N, the width of the array element w, the pitch of the array element p, the height of the array element h, the curvature of the array element R, and the width of the array W.

[0028] In addition to the above size parameters, it also supports customizing the frequency response of the ultrasonic transducer, including the center frequency and the bandwidth. Here, the bandwidth is the relative bandwidth, that is, the ratio of the difference between the upper cut-off frequency and the lower cut-off frequency to the center frequency. The upper cut-off frequency and the lower cut-off frequency are the highest and lowest frequencies corresponding to the intensity dropping to half of the center frequency intensity respectively. Ideally, the ultrasonic signal is a δ function in the time domain with an infinite spectrum, while in reality, the spectrum of the ultrasonic transducer is of finite width. Therefore, in simulated imaging, it is also necessary to emit ultrasonic signals with a finite bandwidth. Figure 4 , which are the frequency-domain spectrum and time-domain spectrum under the frequency response with a center frequency of 5.2 MHz and a relative bandwidth of 60%. Since the ultrasonic transducer needs to both emit and receive ultrasonic signals, it needs to experience two frequency responses, namely the one-way response and the two-way response. The specific implementation steps are to perform band-pass filtering on the ideal impulse response, and the filter uses a first-order Butterworth band-pass filter generated by the defined center frequency and bandwidth. In addition, it also supports defining the arrangement mode of the ultrasonic transducer, which is specifically realized by defining the center coordinates and orientations of each array element. This model supports the simulation of ultrasonic imaging for any array, such as linear arrays, concave arrays, convex arrays, and annular arrays, etc.

[0029] Link 2, ultrasonic signal generation: Ultrasonic signal generation includes two parts: signal transmission and signal acquisition. Four parameters need to be determined for the transmission part: transmission waveform, transmission apodization coefficient, directivity function, and transmission delay. First, the transmission waveform is generated according to the defined frequency response of the ultrasonic transducer, that is, the Figure 4 time-domain signal of the one-way response in. Secondly, the transmission apodization coefficient is the weight coefficient of each transmitting array element, and each array element can be independently controlled. The range of the transmission apodization coefficient is from 0 to 1. 0 means not emitting ultrasonic signals, 1 means emitting ultrasonic signals with the highest intensity, and the remaining values mean emitting ultrasonic signals with different intensities. Then, the directivity function of the array element describes the contribution of the array element to the sound field intensity at different directions, which is determined by the width of the array element and the center frequency, and its expression is: ; In the formula, represents the angle between the specified direction and the normal direction of the array element, represents the wave number of the ultrasonic wave emitted by the transmitting array element, represents half of the array element width, ranges from 0 to 1. Finally, the transmission delay is used to control the wavefront shape of the ultrasonic wave emitted by the transmitting array, such as plane wave, focused spherical wave, and divergent spherical wave.

[0030] After determining these parameters, the shock wave signal at any scattering point can be obtained, specifically as shown in (a) of Figure 5 . The time-domain acoustic wave signal excited by each transmitting array element at the target scattering point is denoted as the shock wave signal, and the calculation method of the shock wave signal is: the transmitted waveform is multiplied by the transmission apodization coefficient of this array element, multiplied by the directivity value at this scattering point, and then multiplied by the attenuation degree of the ultrasonic wave. The attenuation degree of the acoustic wave is determined according to the propagation distance from the transmitting array element to the scattering point. The attenuation of ultrasonic waves in tissues usually follows the exponential attenuation law, and the attenuation function is: ; In the formula, is the initial amplitude, is the attenuation coefficient of the medium, is the frequency of the ultrasonic wave, is the frequency exponent, is the propagation distance of the ultrasonic wave. Then, according to the transmission delay and the propagation distance, the time when the ultrasonic wave reaches this scattering point is determined. Finally, assuming that a total of N transmitting array elements are activated, the time-domain acoustic wave signals excited by all transmitting array elements (numbered #1 to #N) at this scattering point are linearly superimposed to obtain the total shock wave signal of the target scattering point in this transmission sequence. The total shock wave signal is then multiplied by the reflectivity of the scattering point to obtain the initial echo signal of this scattering point.

[0031] For the receiving part, three parameters need to be determined: the received waveform, the receiving apodization coefficient, and the directivity function. The calculation method of the echo signal reaching the receiving array element is: the initial echo signal of the scattering point is applied with the ultrasonic transducer frequency response again, multiplied by the receiving apodization coefficient and the directivity function of the receiving array element, and then multiplied by the attenuation degree of the ultrasonic wave. Then, the attenuation degree of the ultrasonic wave is determined according to the propagation distance from the scattering point to the receiving array element, and then according to the propagation distance from the scattering point to the receiving array element plus the propagation distance from the transmitting array element to the scattering point, the time when the initial echo signal reaches the receiving array element is determined. The time is used to describe the position of the ultrasonic wave in the time domain. Assuming that there are a total of M scattering points (numbered #1 to #M), the time-domain signals of the initial echo signals of all scattering points reaching the receiving array element are linearly superimposed to obtain the total echo signal collected by this receiving channel. The same applies to the other receiving channels, as shown in (b) of Figure 5 .

[0032] Step 3, ultrasonic signal storage: The total echo signal collected is encoded and stored according to the dimensions. The first dimension is the sampling depth. Assuming there are n transmit sequences in total, and 1024 sampling points are collected in each transmit sequence, is n×1024. And the data collected in the first transmit sequence is placed in rows 1 to 1024, the data collected in the second transmit sequence is placed in rows 1025 to 2048, and so on. The second dimension is the number of receiving channels, and the third dimension is the number of frames collected. Each frame can be reconstructed into a B-mode ultrasound image, frames can be reconstructed into B-mode ultrasound images. In addition, the transmit and receive modes of each sequence need to be stored and recorded for subsequent beamforming. Finally, the three-dimensional array and sequence information are packaged and stored in memory or on disk for subsequent signal processing.

[0033] The second step is image reconstruction, also known as the reverse process. First, time gain compensation is performed on the total echo signal to compensate for the acoustic wave attenuation during propagation. Secondly, band-pass filtering is performed on the signal after time gain compensation to filter out interference signals outside the bandwidth. Then, Hilbert transform is performed on the signal of each channel after band-pass filtering for envelope detection. Next, delay focusing is performed on the complex signal, calculating the time from the wavefront at zero time to the pixel point to be reconstructed and then to the receiving array element, respectively indexing the intensities at the corresponding times in the original signal and the transformed signal, and using them as the real and imaginary parts of the pixel point to be reconstructed. Finally, the real part signal and imaginary part signal intensities indexed by each receiving array element during each transmit are added together respectively for beamforming, and the complex value of the pixel point is obtained. Then, taking the modulus of this complex number, the pixel value of the pixel point is obtained. The same applies to the remaining pixel points to be reconstructed. After completing the reconstruction of all pixel points to be reconstructed, the B-mode ultrasound image can be obtained.

[0034] The third step is to calculate the spatial resolution. The quantization method of the spatial resolution is the full width at half maximum (FWHM) of the point spread function of the imaging system. The point spread function refers to the response of an imaging system to an ideal pulse. An ideal scatterer is set in the imaging area, and simulated imaging is performed by the methods of the first two steps. The B-mode ultrasound image of the ideal scatterer obtained is the system point spread function. Then, a profile line of the point spread function is taken along the preset direction, and the value of its FWHM is the value of the spatial resolution in that direction. In B-mode ultrasound imaging, the lateral resolution and axial resolution are usually of concern. The choice of the preset direction depends on which direction of the resolution is to be studied. The resolutions in different directions of space usually have different variation laws. The choice of the preset direction does not affect the resolution quantization result of the imaging system itself. The resolution quantization result of the imaging system itself only depends on the parameters of the imaging system itself. The FWHM of the profile line refers to the spatial span corresponding to the profile line descending to half from the highest point to the left and right sides.

[0035] By changing the parameters in the system, such as the geometry of the array, the frequency response of the ultrasonic transducer, the shape of the array elements, etc., the influence mechanism of these parameters on the spatial resolution can be studied. Two embodiments will be introduced below: (1) Study the influence mechanism of the array geometry on the spatial resolution; (2) Study the influence mechanism of the central frequency of the ultrasonic transducer on the spatial resolution.

[0036] Embodiment 1: By setting ultrasonic transducers with different geometries, such as arc arrays with different angles, the influence of the viewing angle of the ultrasonic transducer on the spatial resolution can be studied. The angles of the arc arrays are set to 360°, 270°, 180°, 135°, 90°, 45°, 22.5°, and 11.25° respectively. Point spread function modeling is performed on these ultrasonic transducers with different shapes, and their spatial resolutions are calculated. The point spread modeling results are as shown in Figure 6 (a). The white dashed circular arcs therein are only used for the schematic of the viewing angle of the arc array ultrasonic transducer and do not represent the actual size of the array diameter. The actual diameter of the ultrasonic transducer array is 80 mm. Figure 6 In (a), the horizontal direction is defined as the lateral direction, and the vertical direction is defined as the axial direction. From Figure 6 (a), it can be seen that as the viewing angle of the ultrasonic transducer decreases, both the lateral and axial resolutions of the imaging system show obvious degradation, and the degradation of the lateral resolution is particularly significant. Especially when the angle is less than 135°, the spatial resolution of the imaging system degrades rapidly. Quantitative calculation of the spatial resolution in Figure 6 (a) is performed, and the results are as shown in Figure 6As shown in (b), it can be seen that the influence laws of the viewing angle change on the lateral and axial resolutions are different. Specifically, when the viewing angle of the ultrasonic transducer gradually decreases from 360° to 135°, the axial resolution rapidly degrades from 80 μm to 242 μm (a significant increase of up to 202%), while the lateral resolution only slightly degrades from 80 μm to 130 μm (an increase of 62%), indicating that the decrease in the viewing angle of the ultrasonic transducer mainly affects the axial resolution in this stage. When the viewing angle is further reduced from 135° to 12.5°, the axial resolution tends to be stable, while the lateral resolution surges from 130 μm to 1491 μm (an increase of more than 10 times), especially showing an exponential degradation trend when the viewing angle is less than 45°. From this set of examples, it can be concluded that the viewing angle of the ultrasonic transducer is an important factor affecting the spatial resolution.

[0037] Example Two: By changing the central frequency of the ultrasonic transducer while keeping other parameters unchanged, the influence of the central frequency of the ultrasonic transducer on the spatial resolution can be studied. Based on the ATL L7-4 linear array ultrasonic transducer (Philips Healthcare), point spread function modeling is carried out. The central frequencies of the ultrasonic transducer are set to 3.5 MHz, 4.5 MHz, 5.2 MHz, 6.3 MHz, 7.8 MHz, 9.0 MHz, and 10.4 MHz respectively. The modeling results of the point spread function at different central frequencies are as Figure 7 shown in (a). As the central frequency increases, the point spread function significantly becomes smaller. For Figure 7 the spatial resolution of (a), quantitative calculation is carried out, and the results are as Figure 7 shown in (b). Both the axial resolution and the lateral resolution are inversely proportional to the central frequency. If the central frequency is converted to the wavelength, the corresponding wavelengths for the above central frequencies are 440 μm, 342 μm, 296 μm, 244 μm, 197 μm, 171 μm, and 148 μm respectively. The relationship between the axial resolution and the wavelength can be approximately written as R A = 1.1λ; the relationship between the lateral resolution and the wavelength can be approximately written as R L = 1.3λ, that is, for the currently set imaging system, the axial resolution is approximately 1.1 times the wavelength, and the lateral resolution is approximately 1.3 times the wavelength. For imaging systems with different parameters, the proportionality coefficient will be different, but it can all be written in the form of R = k·λ. From this set of examples, it can be concluded that both the lateral and axial resolutions are directly proportional to the central frequency and directly proportional to the wavelength.

[0038] The present invention proposes a general spatial resolution research model based on numerical simulation, which can quantify the spatial resolution of any ultrasonic imaging system, and proposes a single-element signal transmission and reception model weighted by three items of directivity function, apodization coefficient and sound attenuation function, as well as an ultrasonic signal generation model of multi-elements and multi-scattering points based on the principle of linear superposition.

[0039] It should be understood that although the steps in the flowchart of the accompanying drawings of the specification are shown in sequence according to the indication of the arrows, these steps are not necessarily executed in the order indicated by the arrows. Unless there is a clear indication in this article, there is no strict order restriction for the execution of these steps, and these steps can be executed in other orders. Moreover, at least a part of the steps in the flowchart of the accompanying drawings of the specification may include multiple steps or multiple stages. These steps or stages are not necessarily executed at the same moment, but can be executed at different moments. The execution order of these steps or stages is not necessarily sequential, but can be executed alternately or alternately with at least a part of other steps or steps or stages in other steps.

[0040] Based on the description of the above method embodiments, the present invention also provides a system. The system may be a system such as software (application), module, component, server, client, etc. that uses the method described in the embodiments of this specification and combines necessary implementation hardware. Based on the same innovative concept, the systems in one or more embodiments provided by the embodiments of the present disclosure are as described in the following embodiments. Since the implementation solutions for the system to solve problems are similar to the method, the implementation of the specific system in the embodiments of this specification may refer to the implementation of the foregoing method, and the repeated parts will not be described again. As used hereinafter, the term "module" or "modular" is a combination of software and / or hardware that can implement a predetermined function. Although the systems described in the following embodiments are preferably implemented in software, implementation in hardware, or a combination of software and hardware is also possible and contemplated.

[0041] Specifically, a general B-ultrasound imaging spatial resolution modeling system in the present invention includes: A parameter configuration module: used to configure ultrasonic transducer parameters; the ultrasonic transducer parameters include one or more of the number of elements in the array, element width, element spacing, element height, element curvature R, array width W, array arrangement method, and ultrasonic transducer frequency response; A signal transmission module: calculates the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point through the transmitted waveform, transmit apodization coefficient, directivity function, and transmit delay, and superimposes the shock wave signals of all transmitting elements to generate a total shock wave signal; multiplies the total shock wave signal by the reflectivity of the scattering point to obtain the initial echo signal of the scattering point; Signal receiving module: Calculate the echo signal after the initial echo signal of a single scatterer arrives at the receiving array element through the received waveform, receive apodization coefficient, directivity function, and receive delay; linearly superimpose the echo signals of the initial echo signals of all scatterers arriving at the current receiving array element to obtain the total echo signal of the current receiving array element, and further obtain the total echo signals of all receiving array elements; Image reconstruction module: Perform image reconstruction on the total echo signal and output a B-mode ultrasound image; Resolution quantization module: Set any point in the imaging area as an ideal scatterer, obtain the point spread function by calculating the B-mode ultrasound image of the ideal scatterer, take the profile line of the point spread function along the preset direction, and the full width at half maximum value of the profile line is the spatial resolution quantization result in this direction; Modeling module: By changing the ultrasonic transducer parameters and calculating the corresponding spatial resolution quantization results, realize the modeling of the influence mechanism of different parameters on the spatial resolution.

[0042] The technical features of the above embodiments can be combined arbitrarily. For the sake of brevity of description, not all possible combinations of the technical features in the above embodiments are described. However, as long as there is no contradiction in the combination of these technical features, it should be considered as the scope described in this specification.

[0043] For those skilled in the art, it is obvious that the present invention is not limited to the details of the above exemplary embodiments, and without departing from the spirit or basic characteristics of the present invention, the present invention can be implemented in other specific forms. Therefore, from any point of view, the embodiments should be regarded as exemplary and non-limiting. The scope of the present invention is defined by the appended claims rather than the above description. Therefore, it is intended to include all changes falling within the meaning and scope of the equivalent elements of the claims in the present invention, and any reference signs in the claims should not be regarded as limiting the claims involved.

[0044] In addition, it should be understood that although this specification is described according to the embodiments, not every embodiment only contains an independent technical solution. This narrative way of the specification is only for the sake of clarity. Those skilled in the art should regard the specification as a whole, and the technical solutions in each embodiment can also be appropriately combined to form other embodiments that can be understood by those skilled in the art.

Claims

1. A general method for modeling the spatial resolution of B-mode ultrasound imaging, characterized in that, Including: Configuring ultrasonic transducer parameters; Calculating the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point through the transmitting waveform, transmitting apodization coefficient, directivity function, and transmitting delay, and superimposing the shock wave signals of all transmitting elements to generate a total shock wave signal; multiplying the total shock wave signal by the reflectivity of the scattering point to obtain the initial echo signal of the scattering point; Calculating the echo signal after the initial echo signal of a single scattering point reaches the receiving element through the receiving waveform, receiving apodization coefficient, directivity function, and receiving delay; linearly superimposing the echo signals of the initial echo signals of all scattering points reaching the current receiving element to obtain the total echo signal of the current receiving element, and further obtaining the total echo signals of all receiving elements; Performing image reconstruction on the total echo signal and outputting a B-mode ultrasound image; Setting any point within the imaging region as an ideal scattering point, obtaining the point spread function by calculating the B-mode ultrasound image of the ideal scattering point, taking the profile line of the point spread function along a preset direction, and the full width at half maximum value of the profile line is the quantization result of the spatial resolution in this direction; By changing the ultrasonic transducer parameters and calculating the corresponding spatial resolution quantization results, a model of the influence mechanism of different parameters on the spatial resolution is realized.

2. The general B-ultrasound imaging spatial resolution modeling method according to claim 1, wherein The configurable ultrasonic transducer parameters include one or more of the number of elements in the array, element width, element spacing, element height, element curvature, array width, array arrangement, and the frequency response of the ultrasonic transducer; the frequency response includes the center frequency and the relative bandwidth, and the relative bandwidth is the ratio of the difference between the upper cut-off frequency and the lower cut-off frequency to the center frequency.

3. A general B-ultrasound imaging spatial resolution modeling method according to claim 1, characterized in that The calculating the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point through the transmitting waveform, transmitting apodization coefficient, directivity function, and transmitting delay specifically includes: The calculation method of the shock wave signal of each transmitting element at the scattering point is: multiplying the transmitting waveform by the transmitting apodization coefficient of the transmitting element, multiplying by the directivity value of the directivity function at the scattering point, and then multiplying by the attenuation function of the ultrasonic wave; Among them, the attenuation function of the ultrasonic wave is determined according to the propagation distance from the element to the scattering point; Determining the position of the shock wave signal in the time domain through the transmitting delay.

4. A general B-ultrasound imaging spatial resolution modeling method according to claim 3, characterized in that, The transmitting waveform is generated according to the frequency response of the ultrasonic transducer, and the transmitting apodization coefficient is the weight coefficient of each transmitting element; the directivity function is determined by the element width and the center frequency.

5. A general B-ultrasound imaging spatial resolution modeling method according to claim 1, characterized in that, The calculating the echo signal after the initial echo signal of a single scattering point reaches the receiving element through the receiving waveform, receiving apodization coefficient, directivity function, and receiving delay specifically includes: The calculation method of the echo signal reaching the receiving element is: applying the frequency response of the ultrasonic transducer to the initial echo signal of the scattering point by the receiving waveform, multiplying by the receiving apodization coefficient of the receiving element, multiplying by the directivity value of the directivity function at the scattering point, and then multiplying by the attenuation function of the ultrasonic wave; Among them, the attenuation function of the ultrasonic wave is determined according to the propagation distance from the scattering point to the receiving element; Determining the position of the echo signal in the time domain through the receiving delay.

6. A general B-ultrasound imaging spatial resolution modeling method according to claim 1, characterized in that The performing image reconstruction on the total echo signal and outputting a B-mode ultrasound image specifically includes: Perform time gain compensation, band-pass filtering, Hilbert transform envelope detection, delay focusing beam synthesis, and logarithmic compression on the total echo signal in sequence, and output a B-mode ultrasound image.

7. A general B-mode ultrasound imaging spatial resolution modeling method according to claim 1, characterized in that The method for obtaining the point spread function by calculating the B-mode ultrasound image of an ideal scatterer specifically includes: Denote the pixel value of any pixel point in the B-mode ultrasound image corresponding to the ideal scatterer as the function f(x, y), where x and y are the row number and column number of the pixel point respectively, and the function f(x, y) is the point spread function.

8. A system using the general B-ultrasound imaging spatial resolution modeling method according to any one of claims 1 to 7, characterized in that, It includes: Parameter configuration module: used to configure ultrasonic transducer parameters; the ultrasonic transducer parameters include one or more of the number of elements in the array, element width, element spacing, element height, element curvature, array width, array arrangement method, and ultrasonic transducer frequency response; Signal transmission module: calculate the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scatterer through the transmission waveform, transmission apodization coefficient, directivity function, and transmission delay, and superimpose the shock wave signals of all transmitting elements to generate a total shock wave signal; multiply the total shock wave signal by the reflectivity of the scatterer to obtain the initial echo signal of the scatterer; Signal reception module: calculate the echo signal after the initial echo signal of a single scatterer reaches the receiving element through the reception waveform, reception apodization coefficient, directivity function, and reception delay; linearly superimpose the echo signals of the initial echo signals of all scatterers reaching the current receiving element to obtain the total echo signal of the current receiving element, and further obtain the total echo signals of all receiving elements; Image reconstruction module: perform image reconstruction on the total echo signal and output a B-mode ultrasound image; Resolution quantization module: set any point in the imaging area as an ideal scatterer, obtain the point spread function by calculating the B-mode ultrasound image of the ideal scatterer, take the profile line of the point spread function along a preset direction, and the full width at half maximum value of the profile line is the spatial resolution quantization result in this direction; Modeling module: model the influence mechanism of different parameters on the spatial resolution by changing the ultrasonic transducer parameters and calculating the corresponding spatial resolution quantization results.

9. The system according to claim 8, wherein The method for calculating the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scatterer through the transmission waveform, transmission apodization coefficient, directivity function, and transmission delay specifically includes: The calculation method of the shock wave signal of each transmitting element at the scatterer is: multiply the transmission waveform by the transmission apodization coefficient of the transmitting element, multiply by the directivity value of the directivity function at the scatterer, and then multiply by the attenuation function of the ultrasonic wave; Wherein, the attenuation function of the ultrasonic wave is determined according to the propagation distance from the element to the scatterer; Determine the position of the shock wave signal in the time domain through the transmission delay.

10. The system according to claim 8, wherein The method for calculating the echo signal after the initial echo signal of a single scatterer reaches the receiving element through the reception waveform, reception apodization coefficient, directivity function, and reception delay specifically includes: The calculation method of the echo signal reaching the receiving element is: apply the ultrasonic transducer frequency response to the initial echo signal of the scatterer by the reception waveform, multiply by the reception apodization coefficient of the receiving element, multiply by the directivity value of the directivity function at the scatterer, and then multiply by the attenuation function of the ultrasonic wave; Among them, the attenuation function of the ultrasonic wave is determined according to the propagation distance from the scattering point to the receiving array element; The position of the echo signal in the time domain is determined by the reception delay.

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