Universal B-ultrasound imaging spatial resolution modeling method and system

Abstract

The present invention relates to the field of ultrasonic medical imaging technology, and discloses a general B-ultrasound imaging spatial resolution modeling method and system; the method includes: configuring ultrasonic transducer parameters; calculating the shock wave signal of a single transmitting array element at any specific scattering point, superimposing all shock wave signals to generate a total shock wave signal, multiplying by the reflectivity of the scattering point to obtain an initial echo signal; calculating the echo signal after reaching the receiving array element; linearly superimposing all echo signals to obtain a total echo signal; performing image reconstruction on the total echo signal to output a B-ultrasound image; obtaining a spatial resolution quantification result through a point spread function; changing the ultrasonic transducer parameters, and calculating the spatial resolution quantification result, to achieve modeling of the mechanism by which different parameters affect spatial resolution. The present invention can support modeling of ultrasonic transducers with arbitrary array arrangements, arbitrary array element orientations, and arbitrary transmitting waveforms, and arbitrary pulse echo sequences, and has a high degree of freedom.

Description

Technical Field

[0001] The present invention relates to the field of ultrasonic medical imaging technology, and in particular to a universal B-ultrasound imaging spatial resolution modeling method and system. Background Art

[0002] B-ultrasound imaging is a radiation-free, non-invasive, and highly real-time medical imaging technology. It reconstructs internal structural information by controlling an ultrasonic transducer to transmit ultrasound waves into a living organism and collect the echo signals. Spatial resolution is a key metric for B-ultrasound imaging, directly impacting diagnostic efficiency and accuracy. Therefore, improving spatial resolution is a key goal in the design and development of B-ultrasound imaging systems.

[0003] For B-ultrasound imaging, numerous factors in the imaging chain, such as array geometry, ultrasonic transducer frequency response, and pulse-echo sequence, affect spatial resolution. Currently, there is a lack of a unified model to study the mechanisms by which these factors influence spatial resolution. Therefore, developing a universal model for studying spatial resolution is of great significance.

[0004] Prior art models exist for studying spatial resolution based on the Field II toolkit, simulating specific imaging scenarios. For example, a model for studying the impact of ultrasonic transducer factors on spatial resolution based on plane-wave imaging is available. This model uses the Field II ultrasound simulation toolkit to simulate imaging of a plane wave sequence from a linear array ultrasonic transducer. By varying ultrasonic transducer parameters such as center frequency, bandwidth, and array width, the impact of these system parameters on spatial resolution is studied. 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 existing art, there are also models for studying spatial resolution based on analytical formulas. These models use the Lippmann-Schwinger equation to establish an analytical expression for the imaging system's point spread function, and study spatial resolution based on the point spread function. The impact of these system parameters on spatial resolution can be studied by varying the number of ultrasonic transducer elements, array width, and bandwidth. Reference: ROQUETTE L, SIMEONI M, HURLEY P, et al. On ananalytical, spatially-varying, point-spread-function[C] / / IEEE International Ultrasonics Symposium (IUS), Washington, USA: IEEE, 2017: 1-4.

[0006] The above-mentioned research model on the influence of ultrasonic transducer factors on spatial resolution based on plane wave imaging has many limitations. First, spatial resolution can only be studied based on plane wave sequences. However, B-ultrasound imaging also has a variety of pulse echo sequences, such as focused line scanning imaging, coherent plane wave composite imaging, and synthetic aperture imaging. Spatial resolution research based on other imaging sequences is also extremely important. Second, only the influence mechanism of ultrasonic transducer parameters on spatial resolution can be studied, and the influence mechanism of other system parameters, such as sequence parameters and beamforming algorithm parameters, cannot be studied. Third, only simple ultrasonic transducer arrays, such as linear arrays, can be simulated for imaging, 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 universal B-ultrasound imaging spatial resolution modeling method and system, which can be applied in various scenarios and support the study of the influence mechanism of multiple system parameters on spatial resolution; it can guide the optimization of system parameters in the B-ultrasound imaging system to improve the spatial resolution of imaging.

[0008] In order to solve the above technical problems, the present invention adopts the following technical solutions:

[0009] In a first aspect, the present invention provides a general B-ultrasound imaging spatial resolution modeling method, comprising:

[0010] Configure ultrasonic transducer parameters;

[0011] Calculating the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point using the transmitting waveform, transmitting apodization coefficient, directivity function, and transmitting delay, superimposing the shock wave signals of all transmitting elements to generate a total shock wave signal; and multiplying the total shock wave signal by the reflectivity of the scattering point to obtain the initial echo signal of the scattering point;

[0012] The echo signal after the initial echo signal of a single scattering point reaches the receiving array element is calculated using the received waveform, the received apodization coefficient, the directivity function, and the received delay. The echo signals of the initial echo signals of all scattering points reaching the current receiving array element are linearly superimposed to obtain the total echo signal of the current receiving array element, and then the total echo signal of all receiving array elements is obtained.

[0013] Perform image reconstruction on the total echo signal and output a B-ultrasound image;

[0014] An arbitrary point in the imaging area is set as an ideal scattering point. The point spread function is obtained by calculating the B-ultrasound image of the ideal scattering point. The profile line of the point spread function is taken along the preset direction. The half-maximum full width of the profile line is the quantitative result of the spatial resolution in that direction.

[0015] By changing the ultrasonic transducer parameters and calculating the corresponding spatial resolution quantification results, the mechanism of the influence of different parameters on spatial resolution can be modeled.

[0016] In one embodiment, the configurable ultrasonic transducer parameters include one or more of the number of array elements in the array, array element width, array element spacing, array element height, array element curvature, array width, array arrangement, and ultrasonic transducer frequency response; the frequency response includes a center frequency and a relative bandwidth, and the relative bandwidth is the ratio of the difference between the upper cutoff frequency and the lower cutoff frequency to the center frequency.

[0017] In one embodiment, the calculation of the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point by using the transmitting waveform, the transmitting apodization coefficient, the directivity function, and the transmitting delay specifically includes:

[0018] The shock wave signal of each transmitting array element at the scattering point is calculated as follows: the transmitting waveform is multiplied by the transmitting apodization coefficient of the transmitting array 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;

[0019] wherein the attenuation function of the ultrasonic wave is determined according to the propagation distance from the array element to the scattering point;

[0020] The position of the shock wave signal in the time domain is determined by transmitting delay.

[0021] In one embodiment, the transmit waveform is generated according to the frequency response of the ultrasonic transducer, the transmit apodization coefficient is a weight coefficient of each transmit array element, and the directivity function is determined by the array element width and the center frequency.

[0022] In one embodiment, calculating the echo signal after the initial echo signal of a single scattering point reaches the receiving array element by using the received waveform, the received apodization coefficient, the directivity function, and the received delay specifically includes:

[0023] The echo signal reaching the receiving element is calculated by applying the ultrasonic transducer frequency response to the initial echo signal at the scattering point, multiplying it by the receiving apodization coefficient of the receiving element, multiplying it by the directivity value of the directivity function at the scattering point, and then multiplying it by the attenuation function of the ultrasonic wave.

[0024] wherein, the attenuation function of the ultrasonic wave is determined according to the propagation distance from the scattering point to the receiving array element;

[0025] The position of the echo signal in the time domain is determined by receiving the delay.

[0026] In one embodiment, performing image reconstruction on the total echo signal to output a B-ultrasound image specifically includes:

[0027] The total echo signal is sequentially subjected to time gain compensation, bandpass filtering, Hilbert transform envelope detection, delayed focusing beamforming and logarithmic compression to output a B-ultrasound image.

[0028] In one embodiment, obtaining a point spread function by calculating an ideal scattering point on a B-ultrasound image specifically includes:

[0029] The pixel value of any pixel point of the B-ultrasound image corresponding to the ideal scattering point is recorded as a function f(x, y), where x and y are the row number and column number of the pixel point respectively. The function f(x, y) is the point spread function.

[0030] In a second aspect, the present invention provides a general B-ultrasound imaging spatial resolution modeling system, comprising:

[0031] Parameter configuration module: used to configure ultrasonic transducer parameters; ultrasonic transducer parameters include one or more of the number of array elements, array element width, array element spacing, array element height, array element curvature, array width, array arrangement, and ultrasonic transducer frequency response;

[0032] Signal transmission module: Calculates the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point through the transmission waveform, transmission tracking coefficient, directivity function and transmission delay, 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;

[0033] Signal receiving module: Calculates the echo signal after the initial echo signal of a single scattering point reaches the receiving array element by receiving the waveform, receiving apodization coefficient, directivity function and receiving delay. Linearly superimposes the echo signals of the initial echo signals of all scattering points reaching the current receiving array element to obtain the total echo signal of the current receiving array element, and then obtains the total echo signal of all receiving array elements.

[0034] Image reconstruction module: performs image reconstruction on the total echo signal and outputs B-ultrasound image;

[0035] Resolution quantification module: Set any point in the imaging area as an ideal scattering point, calculate the point spread function from the B-ultrasound image of the ideal scattering point, and take the profile line of the point spread function along the preset direction. The full width at half maximum of the profile line is the spatial resolution quantification result in that direction;

[0036] Modeling module: By changing the ultrasonic transducer parameters and calculating the corresponding spatial resolution quantification results, the mechanism of the influence of different parameters on spatial resolution is modeled.

[0037] In one embodiment, the calculation of the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point by using the transmitting waveform, the transmitting apodization coefficient, the directivity function, and the transmitting delay specifically includes:

[0038] The shock wave signal of each transmitting array element at the scattering point is calculated as follows: the transmitting waveform is multiplied by the transmitting apodization coefficient of the transmitting array 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;

[0039] wherein the attenuation function of the ultrasonic wave is determined according to the propagation distance from the array element to the scattering point;

[0040] The position of the shock wave signal in the time domain is determined by transmitting delay.

[0041] In one embodiment, calculating the echo signal after the initial echo signal of a single scattering point reaches the receiving array element by using the received waveform, the received apodization coefficient, the directivity function, and the received delay specifically includes:

[0042] The echo signal reaching the receiving element is calculated by applying the ultrasonic transducer frequency response to the initial echo signal at the scattering point, multiplying it by the receiving apodization coefficient of the receiving element, multiplying it by the directivity value of the directivity function at the scattering point, and then multiplying it by the attenuation function of the ultrasonic wave.

[0043] wherein, the attenuation function of the ultrasonic wave is determined according to the propagation distance from the scattering point to the receiving array element;

[0044] The position of the echo signal in the time domain is determined by receiving the delay.

[0045] Compared with the prior art, the beneficial technical effects of the present invention are:

[0046] For the generation of ultrasonic signals, the signal acquisition model proposed in the present invention can support the modeling of ultrasonic transducers with arbitrary array arrangements, arbitrary array element orientations, arbitrary emission waveforms, and arbitrary pulse-echo sequences, and has a high degree of freedom.

[0047] Compared to spatial resolution research models based on analytical formulas, the numerical simulation-based research model of the present invention has lower complexity and computational effort. Compared to spatial resolution research models based on the Field II toolkit, the present invention offers greater flexibility, supporting the modeling of any ultrasonic transducer and any pulse-echo sequence. BRIEF DESCRIPTION OF THE DRAWINGS

[0048] Figure 1 is a flow chart of a method in an embodiment of the present invention;

[0049] Figure 2 A specific flow chart for digital modeling of a B-ultrasound imaging system in an embodiment of the present invention;

[0050] Figure 3 This is a schematic diagram of some parameters supported by customization of the linear array ultrasonic transducer in an embodiment of the present invention;

[0051] Figure 4 The frequency domain spectrum and time domain spectrum of the transmitted ultrasound at the set center frequency and bandwidth in the embodiment of the present invention;

[0052] Figure 5 Schematic diagram of the principle of an ultrasonic signal generation model based on linear superposition in an embodiment of the present invention;

[0053] Figure 6 The point spread function modeling results and spatial resolution quantification results of ultrasonic transducers with different geometric shapes in Example 1;

[0054] Figure 7 These are the point spread function modeling results and spatial resolution quantification results of ultrasonic transducers with different center frequencies in Example 2. DETAILED DESCRIPTION

[0055] A preferred embodiment of the present invention will be described in detail below with reference to the accompanying drawings.

[0056] like Figure 1 As shown, the present invention provides a general B-ultrasound imaging spatial resolution modeling method, comprising the following steps:

[0057] S1, configure ultrasonic transducer parameters;

[0058] S2, calculating the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point by using the transmitting waveform, transmitting apodization coefficient, directivity function, and transmitting delay, superimposing the shock wave signals of all transmitting elements 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 of the scattering point;

[0059] S3, calculating the echo signal after the initial echo signal of a single scattering point reaches the receiving element by using the received waveform, the received apodization coefficient, the directivity function, and the received 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 then obtaining the total echo signal of all receiving elements;

[0060] S4, performing image reconstruction on the total echo signal and outputting a B-ultrasound image;

[0061] S5, setting any point in the imaging area as an ideal scattering point, obtaining a point spread function by calculating the B-ultrasound image of the ideal scattering point, taking a profile line of the point spread function along a preset direction, and the full width at half maximum of the profile line is the quantified result of the spatial resolution in that direction;

[0062] S6, by changing the ultrasonic transducer parameters and calculating the corresponding spatial resolution quantification results, the mechanism of the influence of different parameters on spatial resolution is modeled.

[0063] The present invention supports digital modeling of B-ultrasound imaging systems under various parameter settings. The specific process of modeling is as follows: Figure 2As shown in the figure. First, the ultrasonic transducer sets the ultrasonic waveform to be emitted. The ultrasonic transducer transmits and receives ultrasonic signals according to the set pulse-echo sequence, which includes a transmit sequence and a receive sequence. Then, the original digital ultrasonic signal is encoded, arranged, and stored according to the execution order of the pulse-echo sequence. Next, signal processing is performed. After obtaining the digital ultrasonic signal, since the sound wave attenuates as it propagates in space, intensity compensation based on the sound wave propagation time is required, which is called time gain compensation. Next, the signal is bandpass filtered to eliminate interference from clutter signals and system noise outside the ultrasonic transducer bandwidth. Subsequently, the signal needs to be Hilbert transformed, with the original signal as the real part and the transformed signal as the imaginary part, which is called envelope detection. This is followed by image reconstruction, which calculates the time index of each pixel in each pulse-echo sequence corresponding to the signal, called delay focusing. Based on the calculated delay value, the corresponding intensity value of the real and imaginary parts of the signal is indexed separately, and then weighted summed. This process is called beamforming. Each pixel in the beamformed image is a complex value, and taking the modulus length yields the reconstructed pixel value. To visualize structures within biological tissue where echo signals are weak, the image needs to be compressed, often using logarithmic compression.

[0064] Logarithmic compression first normalizes the image pixel values, then takes the common logarithm of each pixel value and multiplies it by 20. The calculation formula is as follows:

[0065] ;

[0066] Where, is the original pixel value of the i-th pixel in the image, is the maximum original pixel value, is the compressed pixel value of the i-th pixel in the image, expressed in dB. The compressed pixel value range is (-∞, 0], so a lower threshold is selected such that all pixel values ​​below this threshold are equal to the threshold. For example, if -60 is selected as the lower threshold, the image dynamic range is said to be 60 dB. By selecting different lower thresholds, varying degrees of image compression can be achieved. Finally, after obtaining the B-ultrasound image, the image is displayed. By simulating imaging of an ideal scattering point and obtaining an image of the imaging system's point spread function, spatial resolution can be quantified.

[0067] The specific process of the present invention is divided into three steps: ultrasonic signal acquisition, ultrasonic image reconstruction and spatial resolution calculation.

[0068] The first step is ultrasonic signal acquisition, also known as the forward process. This refers to the process of controlling the ultrasonic transducer to emit ultrasonic waves and collect echoes. Digital modeling of point target echo signals is achieved based on a linear model. The core process includes three key steps: ultrasonic transducer definition, ultrasonic signal generation, and ultrasonic signal storage.

[0069] Link 1, Ultrasonic transducer definition: The present invention supports customizing various parameters of ultrasonic transducers. Taking the commonly used linear array ultrasonic transducer as an example, the customizable parameters are as follows: Figure 3 As shown in Figure 1, it includes: number of array elements N, array element width w, array element spacing p, array element height h, array element curvature R and array width W.

[0070] In addition to the aforementioned dimensional parameters, the frequency response of the ultrasonic transducer can also be customized, including the center frequency and bandwidth. The bandwidth here is the relative bandwidth, which is the ratio of the difference between the upper and lower cutoff frequencies to the center frequency. The upper and lower cutoff frequencies are the highest and lowest frequencies corresponding to the intensity dropping to half the center frequency, respectively. Ideally, an ultrasonic signal in the time domain is a delta function with an infinite spectrum. However, in reality, the spectrum of an ultrasonic transducer is finite. Therefore, in analog imaging, it is also necessary to transmit an ultrasonic signal with a finite bandwidth. Figure 4 , which are the frequency domain spectrum and time domain spectrum under the frequency response of the center frequency of 5.2MHz and the relative bandwidth of 60%. Because the ultrasonic transducer needs to both transmit and receive ultrasonic signals, it needs to undergo two frequency responses, one-way response and one-way response. The specific implementation steps are to perform bandpass filtering on the ideal impulse response. The filter uses a first-order Butterworth bandpass filter generated by the defined center frequency and bandwidth. In addition, it also supports the definition of the arrangement of ultrasonic transducers, which is achieved by defining the center coordinates and orientation of each array element. This model supports simulated ultrasonic imaging of arbitrary arrays, such as linear arrays, concave arrays, convex arrays, and circular arrays.

[0071] Step 2, ultrasonic signal generation: Ultrasonic signal generation includes two parts: signal transmission and signal acquisition. The transmission part needs to determine four parameters: transmission waveform, transmission tracking coefficient, directivity function and transmission delay. First, the transmission waveform is generated according to the defined ultrasonic transducer frequency response, which is Figure 4 The corresponding time domain signal for a single pass in the array. Secondly, the transmit apodization coefficient is the weight coefficient of each transmit array element, and each array element can be controlled independently. The transmit apodization coefficient ranges from 0 to 1, where 0 means no ultrasonic signal is transmitted, 1 means the ultrasonic signal with the highest intensity is transmitted, and the remaining values ​​indicate the transmission of ultrasonic signals of varying intensities. Then, the directivity function of the array element describes the contribution of the array element to the sound field intensity in different directions. It is determined by the array element width and center frequency and is expressed as:

[0072] ;

[0073] Where, Represents the angle between the specified direction and the array element normal, It represents the wave number of the ultrasonic wave emitted by the transmitting array element. represents half of the array element width, The range is 0 to 1. Finally, the transmit delay is used to control the wavefront shape of the ultrasonic wave emitted by the transmit array, such as plane wave, focused spherical wave and divergent spherical wave.

[0074] After determining these parameters, the shock wave signal at any scattering point can be obtained, as shown in the following example: Figure 5 As shown in (a) in the figure. The time domain acoustic wave signal excited by each transmitting array element at the target scattering point is recorded as the shock wave signal. The shock wave signal is calculated as follows: the transmitting waveform is multiplied by the transmitting trace coefficient of the array element, multiplied by the directivity value at the scattering point, and then multiplied by the attenuation of the ultrasonic wave. The attenuation of the acoustic wave is determined by the propagation distance from the transmitting array element to the scattering point. The attenuation of ultrasonic waves in tissue usually follows the exponential attenuation law, and the attenuation function is: for:

[0075] ;

[0076] Where, is the initial amplitude, is the attenuation coefficient of the medium, is the frequency of the ultrasound, is the frequency index, is the propagation distance of the ultrasound. The time it takes for the ultrasound to reach the scattering point is then determined based on the transmission delay and propagation distance. Finally, assuming a total of N transmitting elements are activated, the time-domain acoustic wave signals generated by all transmitting elements (numbered #1 to #N) at the scattering point are linearly superimposed to obtain the total shock wave signal at 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 from that scattering point.

[0077] For the receiving part, three parameters need to be determined: receiving waveform, receiving tracking coefficient and directivity function. The echo signal reaching the receiving array element is calculated as follows: apply the ultrasonic transducer frequency response to the initial echo signal of the scattering point again, multiply it by the receiving tracking coefficient and directivity function of the receiving array element, and then multiply it by the attenuation degree of the ultrasonic wave. The attenuation degree of the ultrasonic wave is then determined based on the propagation distance from the scattering point to the receiving array element, and then the time it takes for the initial echo signal to reach the receiving array element is determined based on 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 is used to describe the position of the ultrasonic wave in the time domain. Assume that there are M scattering points in total (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 the receiving channel. The same is true for the remaining receiving channels, such as Figure 5 As shown in (b) in .

[0078] Step 3: Ultrasonic signal storage: The total echo signal collected is stored according to The first dimension is encoded and stored. is the sampling depth. Assuming there are n transmission sequences, each transmission sequence collects 1024 sampling points. The data collected by the first transmission sequence is placed in rows 1 to 1024, the data collected by the second transmission sequence is placed in rows 1025 to 2048, and so on. The third dimension is the number of receiving channels. is the number of frames collected. Each frame can reconstruct a B-ultrasound image. Frames can be reconstructed In addition, the transmission and reception modes of each sequence need to be stored and recorded for subsequent beam synthesis. The three-dimensional array and sequence information are packaged and stored in memory or hard disk for subsequent signal processing.

[0079] The second step is image reconstruction, also known as the inverse process. First, the total echo signal is time-gain compensated to compensate for acoustic attenuation during propagation. Second, the time-gain-compensated signal is bandpass filtered to remove interference signals outside the bandwidth. Then, the signal of each channel after bandpass filtering is Hilbert transformed for envelope detection. Next, the complex signal is time-delayed focused, calculating the time from the wavefront at time zero to the pixel to be reconstructed and then to the receiving element. The intensity of the corresponding time in the original signal and the transformed signal is indexed as the real and imaginary parts of the pixel to be reconstructed, respectively. Finally, the real and imaginary signal intensities indexed by each receiving element during each transmission are summed and beamformed to obtain the complex value of the pixel. The modulus of this complex value is then taken to obtain the pixel value of the pixel. The same process is repeated for the remaining pixels to be reconstructed. Once all pixels to be reconstructed are reconstructed, the ultrasound image is obtained.

[0080] The third step is to calculate spatial resolution. Spatial resolution is quantified using the full width at half maximum (FWHM) of the imaging system's point spread function (PSF). The PSF refers to the imaging system's response to an ideal pulse. An ideal scattering point is set within the imaging area, and simulated imaging is performed using the methods of the first two steps. The resulting ultrasound image of the ideal scattering point is the system's PSF. A profile line of the PSF is then taken along a preset direction, and its FWHM value represents the spatial resolution in that direction. Ultrasound imaging typically focuses on lateral and axial resolution. The choice of preset direction depends on the desired resolution in which direction, and spatial resolution in different directions typically exhibits different variations. The choice of preset direction does not affect the resolution quantification of the imaging system itself, which depends solely on its parameters. The FWHM of the profile line refers to the spatial span corresponding to the point where the profile line descends halfway from its highest point to the left and right sides.

[0081] By varying system parameters such as the array geometry, the frequency response of the ultrasonic transducer, and the shape of the array elements, we can investigate how these parameters influence spatial resolution. Two examples are presented below: (1) investigating how array geometry influences spatial resolution; and (2) investigating how the center frequency of the ultrasonic transducer influences spatial resolution.

[0082] Example 1:

[0083] By setting ultrasonic transducers of different geometric shapes, such as arc arrays of different angles, the effect of the ultrasonic transducer's viewing angle on the spatial resolution can be studied. The angles of the arc array are set to 360°, 270°, 180°, 135°, 90°, 45°, 22.5°, and 11.25°, and the point spread function of these ultrasonic transducers of different shapes is modeled and their spatial resolution is calculated. The point spread modeling results are shown in Figure 2. Figure 6 As shown in (a), the white dotted arc is only used as a schematic diagram of the viewing angle of the arc array ultrasonic transducer and does not represent the actual size of the array diameter. The actual ultrasonic transducer array diameter is 80 mm. Figure 6 In (a), the horizontal direction is defined as the transverse direction and the vertical direction is defined as the axial direction. Figure 6 As can be seen in (a), as the viewing angle of the ultrasonic transducer decreases, the lateral and axial resolutions of the imaging system degrade significantly, with the degradation of the lateral resolution being particularly significant. In particular, when the viewing angle is below 135°, the spatial resolution of the imaging system degrades sharply. Figure 6 The spatial resolution in (a) is quantitatively calculated, and the results are as follows Figure 6 As shown in (b), it can be seen that the effect of viewing angle changes on lateral and axial resolution varies. Specifically, when the ultrasonic transducer viewing angle is gradually reduced from 360° to 135°, the axial resolution degrades sharply from 80μm to 242μm (an increase of 202%), while the lateral resolution only degrades slightly from 80μm to 130μm (an increase of 62%), indicating that the reduction in the ultrasonic transducer viewing angle at this stage primarily affects the axial resolution. When the viewing angle is further reduced from 135° to 12.5°, the axial resolution stabilizes, while the lateral resolution increases sharply from 130μm to 1491μm (an increase of more than 10 times), especially after the viewing angle is less than 45°, showing an exponential degradation trend. This set of examples concludes that the viewing angle of the ultrasonic transducer is an important factor affecting spatial resolution.

[0084] Example 2:

[0085] By changing the center frequency of the ultrasonic transducer and keeping other parameters unchanged, the effect of the center frequency of the ultrasonic transducer on the spatial resolution can be studied. Based on the ATL L7-4 linear array ultrasonic transducer (Philips Healthcare), the point spread function modeling was performed. The center frequency of the ultrasonic transducer was set to 3.5 MHz, 4.5 MHz, 5.2 MHz, 6.3 MHz, 7.8 MHz, 9.0 MHz and 10.4 MHz. The point spread function modeling results at different center frequencies are shown in Figure 2. Figure 7 As shown in (a), as the center frequency increases, the point spread function becomes smaller. Figure 7 The spatial resolution of (a) is quantitatively calculated, and the results are as follows Figure 7 As shown in (b), both the axial resolution and the lateral resolution are inversely proportional to the center frequency. If the center frequency is converted to wavelength, the wavelengths corresponding to the above center frequencies are 440μm, 342μm, 296μm, 244μm, 197μm, 171μm and 148μm respectively. The relationship between axial resolution and wavelength can be approximately written as R A=1.1λ; the relationship between lateral resolution and wavelength can be approximately written as R L =1.3λ, meaning that for the currently configured imaging system, the axial resolution is approximately 1.1 times the wavelength, and the lateral resolution is approximately 1.3 times the wavelength. The proportionality factor varies for imaging systems with different parameters, but can all be expressed as R = k·λ. This set of examples demonstrates that both lateral and axial resolution are directly proportional to the center frequency and wavelength.

[0086] The present invention proposes a universal spatial resolution research model based on numerical simulation, which can quantify the spatial resolution of any ultrasonic imaging system. It also proposes a single-element signal transmission and reception model based on the weighting of directivity function, tracking coefficient and acoustic attenuation function, as well as a multi-element, multi-scattering point ultrasonic signal generation model based on the linear superposition principle.

[0087] It should be understood that although the steps in the flowcharts of the accompanying drawings are shown in sequence as indicated by the arrows, these steps are not necessarily executed in the order indicated by the arrows. Unless otherwise specified herein, there is no strict order restriction on the execution of these steps, and these steps can be executed in other orders. Moreover, at least some of the steps in the flowcharts of the accompanying drawings may include multiple steps or multiple stages, and these steps or stages are not necessarily executed at the same time, but can be executed at different times. The order of execution of these steps or stages is not necessarily sequential, but can be executed in turn or alternately with other steps or at least a portion of the steps or stages in other steps.

[0088] Based on the description of the above method embodiments, the present invention also provides a system. The system can be a system that uses the software (application), module, component, server, client, etc. of the method described in the embodiments of this specification and is combined with the necessary implementation hardware. Based on the same innovative concept, the system in one or more embodiments provided by the embodiments of the present disclosure is as described in the following embodiments. Since the implementation scheme and method for solving the problem of the system are similar, the implementation of the specific system of the embodiments of this specification can refer to the implementation of the aforementioned method, and the repeated parts will not be repeated. As used below, the term "module" or "module" refers to a combination of software and / or hardware that can realize the predetermined function. Although the system described in the following embodiments is preferably implemented in software, the implementation of hardware, or a combination of software and hardware, is also possible and conceived.

[0089] Specifically, a general B-ultrasound imaging spatial resolution modeling system in the present invention includes:

[0090] Parameter configuration module: used to configure ultrasonic transducer parameters; ultrasonic transducer parameters include one or more of the number of array elements, array element width, array element spacing, array element height, array element curvature R, array width W, array arrangement mode, and ultrasonic transducer frequency response;

[0091] Signal transmission module: Calculates the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point through the transmission waveform, transmission tracking coefficient, directivity function and transmission delay, 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;

[0092] Signal receiving module: Calculates the echo signal after the initial echo signal of a single scattering point reaches the receiving array element by receiving the waveform, receiving apodization coefficient, directivity function and receiving delay. Linearly superimposes the echo signals of the initial echo signals of all scattering points reaching the current receiving array element to obtain the total echo signal of the current receiving array element, and then obtains the total echo signal of all receiving array elements.

[0093] Image reconstruction module: performs image reconstruction on the total echo signal and outputs B-ultrasound image;

[0094] Resolution quantification module: Set any point in the imaging area as an ideal scattering point, calculate the point spread function from the B-ultrasound image of the ideal scattering point, and take the profile line of the point spread function along the preset direction. The full width at half maximum of the profile line is the spatial resolution quantification result in that direction;

[0095] Modeling module: By changing the ultrasonic transducer parameters and calculating the corresponding spatial resolution quantification results, the mechanism of the influence of different parameters on spatial resolution is modeled.

[0096] The technical features of the above embodiments can be combined arbitrarily. To make the description concise, 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, they should be considered to be within the scope of this specification.

[0097] It will be apparent to those skilled in the art that the present invention is not limited to the details of the exemplary embodiments described above and that the invention can be embodied in other specific forms without departing from the spirit or essential characteristics of the invention. Therefore, the embodiments should be considered in all respects as illustrative and non-limiting, and the scope of the invention is defined by the appended claims rather than the foregoing description. It is intended that all variations within the meaning and range of equivalents of the claims be embraced herein, and any reference signs in the claims should not be construed as limiting the claims to which they relate.

[0098] In addition, it should be understood that although this specification is described in terms of implementation methods, not every implementation method contains only one independent technical solution. This narrative method of the specification is only for the sake of clarity. Those skilled in the art should regard the specification as a whole. The technical solutions in each embodiment can also be appropriately combined to form other implementation methods that can be understood by those skilled in the art.

Claims

1. A general B-ultrasound imaging spatial resolution modeling method, characterized in that: include: Configure ultrasonic transducer parameters; The shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point is calculated using the transmitting waveform, transmitting apodization coefficient, directivity function, and transmitting delay, and the shock wave signals of all transmitting elements are superimposed to generate a total shock wave signal; the total shock wave signal is multiplied by the reflectivity of the scattering point to obtain the initial echo signal of the scattering point; wherein the time domain acoustic wave signal excited by each transmitting element at the target scattering point is recorded as the shock wave signal; The echo signal after the initial echo signal of a single scattering point reaches the receiving array element is calculated using the received waveform, the received apodization coefficient, the directivity function, and the received delay. The echo signals of the initial echo signals of all scattering points reaching the current receiving array element are linearly superimposed to obtain the total echo signal of the current receiving array element, and then the total echo signal of all receiving array elements is obtained. Perform image reconstruction on the total echo signal and output a B-ultrasound image; An arbitrary point in the imaging area is set as an ideal scattering point. The point spread function is obtained by calculating the B-ultrasound image of the ideal scattering point. The profile line of the point spread function is taken along the preset direction. The half-maximum full width of the profile line is the quantitative result of the spatial resolution in that direction. By changing the ultrasonic transducer parameters and calculating the corresponding spatial resolution quantification results, the mechanism of the influence of different parameters on spatial resolution can be modeled.

2. A general B-ultrasound imaging spatial resolution modeling method according to claim 1, characterized in that: Configurable ultrasonic transducer parameters include one or more of the number of array elements in the array, array element width, array element spacing, array element height, array element curvature, array width, array arrangement, and ultrasonic transducer frequency response; the frequency response includes a center frequency and a relative bandwidth, where the relative bandwidth is the ratio of the difference between the upper cutoff frequency and the lower cutoff frequency to the center frequency.

3. A general B-ultrasound imaging spatial resolution modeling method according to claim 1, characterized in that: The calculation of the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point by using the transmitting waveform, the transmitting apodization coefficient, the directivity function and the transmitting delay specifically includes: The shock wave signal of each transmitting array element at the scattering point is calculated as follows: the transmitting waveform is multiplied by the transmitting apodization coefficient of the transmitting array 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 array element to the scattering point; The position of the shock wave signal in the time domain is determined by transmitting delay.

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

5. A general B-ultrasound imaging spatial resolution modeling method according to claim 1, characterized in that: The calculation of the echo signal after the initial echo signal of a single scattering point reaches the receiving array element by using the receiving waveform, the receiving apodization coefficient, the directivity function and the receiving delay specifically includes: The echo signal reaching the receiving element is calculated by applying the ultrasonic transducer frequency response to the initial echo signal at the scattering point, multiplying it by the receiving apodization coefficient of the receiving element, multiplying it by the directivity value of the directivity function at the scattering point, and then multiplying it 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 array element; The position of the echo signal in the time domain is determined by receiving the 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 to output a B-ultrasound image specifically includes: The total echo signal is sequentially subjected to time gain compensation, bandpass filtering, Hilbert transform envelope detection, delayed focusing beamforming and logarithmic compression to output a B-ultrasound image.

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

8. A system using the universal B-ultrasound imaging spatial resolution modeling method according to any one of claims 1 to 7, characterized in that: include: Parameter configuration module: used to configure ultrasonic transducer parameters; ultrasonic transducer parameters include one or more of the number of array elements, array element width, array element spacing, array element height, array element curvature, array width, array arrangement, and ultrasonic transducer frequency response; Signal transmission module: Calculates the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point by using the transmission waveform, transmission apodization coefficient, directivity function, and transmission 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; wherein, the time domain acoustic wave signal excited by each transmitting element at the target scattering point is recorded as the shock wave signal; Signal receiving module: Calculates the echo signal after the initial echo signal of a single scattering point reaches the receiving array element by receiving the waveform, receiving apodization coefficient, directivity function and receiving delay. Linearly superimposes the echo signals of the initial echo signals of all scattering points reaching the current receiving array element to obtain the total echo signal of the current receiving array element, and then obtains the total echo signal of all receiving array elements. Image reconstruction module: performs image reconstruction on the total echo signal and outputs B-ultrasound image; Resolution quantification module: Set any point in the imaging area as an ideal scattering point, calculate the point spread function from the B-ultrasound image of the ideal scattering point, and take the profile line of the point spread function along the preset direction. The full width at half maximum of the profile line is the spatial resolution quantification result in that direction; Modeling module: By changing the ultrasonic transducer parameters and calculating the corresponding spatial resolution quantification results, the mechanism of the influence of different parameters on spatial resolution is modeled.

9. The system according to claim 8, characterized in that The calculation of the shock wave signal of a single transmitting element of the ultrasonic transducer at any specific scattering point by using the transmitting waveform, the transmitting apodization coefficient, the directivity function and the transmitting delay specifically includes: The shock wave signal of each transmitting array element at the scattering point is calculated as follows: the transmitting waveform is multiplied by the transmitting apodization coefficient of the transmitting array 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 array element to the scattering point; The position of the shock wave signal in the time domain is determined by transmitting delay.

10. The system according to claim 8, wherein: The calculation of the echo signal after the initial echo signal of a single scattering point reaches the receiving array element by using the receiving waveform, the receiving apodization coefficient, the directivity function and the receiving delay specifically includes: The echo signal reaching the receiving element is calculated by applying the ultrasonic transducer frequency response to the initial echo signal at the scattering point, multiplying it by the receiving apodization coefficient of the receiving element, multiplying it by the directivity value of the directivity function at the scattering point, and then multiplying it 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 array element; The position of the echo signal in the time domain is determined by receiving the delay.

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