Ultrasonic imaging system and method based on heart surface, medium and product

Abstract

The invention discloses an ultrasonic imaging system and method based on a heart surface, a medium and a product, the system comprises an ultrasonic detection assembly and an imaging terminal which are in communication connection, an ultrasonic transducer is packaged at the head end of an attached ultrasonic probe in the ultrasonic detection assembly, and an acoustic window on the ultrasonic transducer is planar; the ultrasonic transducer is used for being attached to the surface of the heart or the aorta, and the plane of the ultrasonic transducer is parallel to the plane of the acoustic window; and the imaging terminal is used for generating a sound velocity mask pattern according to the reference ultrasonic image, determining echo time delay data according to the sound velocity mask pattern, synthesizing and reconstructing the radio frequency signal data according to the echo time delay data to obtain an ultrasonic image, performing motion compensation on the image sequence, performing multi-modal fusion on the ultrasonic image subjected to motion compensation, and obtaining a final ultrasonic image. The problem of acoustic interference of anatomical structures such as the chest wall and the esophagus in a traditional echocardiography technology is solved, and real-time and high-quality image support is provided for cardiac surgery.

Description

Technical Field

[0001] This invention relates to the field of echocardiography technology, and more particularly to an ultrasound imaging system, method, medium, and product based on the surface of the heart. Background Technology

[0002] Echocardiography is a non-invasive medical imaging technique that uses high-frequency sound waves to image the heart and its related structures. It is widely used in the diagnosis and treatment of cardiovascular diseases. Currently, the main echocardiographic techniques include transthoracic echocardiography (TTE) and transesophageal echocardiography (TEE). These two techniques each have their advantages and disadvantages and are suitable for different clinical scenarios.

[0003] Transthoracic echocardiography (TTE) is a commonly used non-invasive technique that uses an attached ultrasound probe placed on the chest wall surface for imaging. It offers advantages such as ease of operation and non-invasiveness. However, in cardiac surgery, TTE has the following limitations:

[0004] First, TTE probes suffer from large size and insufficient flexibility. Limited by the size of the incision in open-chest surgery, the size and mechanical design of the TTE probe restrict its flexible operation within the thoracic cavity, a limitation even more pronounced in pediatric cardiac surgery. Furthermore, the TTE probe can only probe from the front of the heart, resulting in the inability to acquire some necessary cardiac cross-sectional images and creating imaging blind spots. For example, the sides and back of the heart cannot be clearly observed, making a comprehensive assessment of cardiac structure difficult. Second, when attempting to directly adhere the TTE probe to the heart surface, insufficient contact can easily create air gaps, leading to decreased image quality. To improve image quality and achieve a closer fit, a certain amount of pressure must be applied to the heart, which can easily affect vital signs such as heart rate and blood pressure during surgery, increasing intraoperative risks. Finally, the TTE probe is a reusable probe that needs to be covered with a sterile bag before use. This operation method not only increases the complexity of use and the risk of postoperative infection, but also reduces the doctor's operating experience and leads to insufficient real-time guidance during the operation. In addition, the mechanical structure design of the TTE probe does not fully consider the surgeon's usage habits and is incompatible with the operation methods of other surgical equipment (such as high-frequency electrosurgical units), making it difficult for the doctor to operate the probe smoothly during the operation.

[0005] Transesophageal echocardiography (TEE) involves inserting a flexible tube with an ultrasound probe into the esophagus to image the heart. It provides high-quality images of the heart and aorta and reduces interference from the ribs and lungs, making it a common choice for intraoperative ultrasound imaging. However, TEE technology also has the following limitations in cardiac surgery:

[0006] First, transesophageal echocardiography (TEE) is an invasive procedure that carries risks of postoperative complications, such as throat injury, esophageal mucosal tears, and gastrointestinal bleeding. This risk may further increase during complex surgeries due to prolonged system use. Furthermore, TEE is contraindicated in patients with anatomical variations in the throat or esophagus, those with digestive diseases, the elderly, and some newborns. Second, TEE can only image from behind the heart, limiting the imaging angle and creating blind spots. For example, due to the distance between the esophagus and the right ventricle, TEE provides insufficient visualization of overall right ventricular function and outflow tract details, and is lacking in detail regarding subtle lesions at the apex and anterior wall. Additionally, probe placement may be affected by anatomical structures such as the esophagus, trachea, and main bronchus, leading to decreased image quality, especially when observing fine cardiac structures, potentially failing to meet the requirements for high-resolution imaging. Finally, TEE equipment is expensive to purchase and maintain, the probe is difficult to operate, and the operator must possess advanced cardiac ultrasound qualifications. Surgeons cannot directly adjust the probe position intraoperatively for real-time guidance, limiting its flexibility in complex surgeries.

[0007] In summary, while transthoracic echocardiography (TTE) and transesophageal echocardiography (TEE) each have their advantages, they also have significant limitations in cardiac surgery. Therefore, given the complex requirements of cardiac surgery, it is of great clinical significance to develop a cardiac ultrasound probe that offers unrestricted exploration angles, high-quality imaging without interference, no risk of complications, and simple and practical operation. Summary of the Invention

[0008] This invention provides an ultrasound imaging system, method, medium, and product based on the surface of the heart to solve the problem of acoustic interference from anatomical structures such as the chest wall and esophagus in traditional echocardiography, thereby improving the quality of ultrasound imaging of the heart structure during cardiac surgery.

[0009] One embodiment of the present invention provides an ultrasound imaging system for the surface of the heart, comprising: an ultrasound detection component and an imaging terminal connected in communication, wherein the ultrasound detection component includes an attached ultrasound probe;

[0010] The attached ultrasound probe has an ultrasound transducer encapsulated at its head end. The acoustic window on the ultrasound transducer is planar and is used to attach to the surface of the heart or aorta. The plane of the ultrasound transducer is parallel to the plane of the acoustic window.

[0011] The imaging terminal is used to receive radio frequency signal data acquired by the attached ultrasound probe, generate a sound velocity mask based on a reference ultrasound image corresponding to the radio frequency signal data, the sound velocity mask containing at least two sound velocity distribution regions, determine echo delay data based on the sound velocity mask, and synthesize and reconstruct an ultrasound image from the radio frequency signal data based on the echo delay data, perform motion compensation on an image sequence composed of multiple ultrasound images, and perform multimodal fusion on the motion-compensated ultrasound image to obtain the final ultrasound image.

[0012] Another embodiment of the present invention provides a cardiac surface-based ultrasound imaging method, applied to the cardiac surface-based ultrasound imaging system described in any embodiment of the present invention, the method comprising:

[0013] Acquire radiofrequency signal data from an attached ultrasound probe on the surface of the heart or aorta;

[0014] A sound velocity mask is generated based on a reference ultrasonic image corresponding to the radio frequency signal data, the sound velocity mask containing at least two sound velocity distribution regions;

[0015] Based on the sound velocity mask, echo delay data is determined, and based on the echo delay data, the radio frequency signal data is synthesized and reconstructed to obtain an ultrasonic image;

[0016] Motion compensation is performed on an image sequence composed of multiple ultrasound images, and multimodal fusion is performed on the motion-compensated ultrasound images to obtain the final ultrasound image.

[0017] According to another embodiment of the present invention, a computer-readable storage medium is provided, the computer-readable storage medium storing computer instructions for causing a processor to execute and implement the cardiac surface-based ultrasound imaging method according to any embodiment of the present invention.

[0018] According to another embodiment of the present invention, a computer program product is provided, including a computer program that, when executed by a processor, implements the cardiac surface-based ultrasound imaging method described in any embodiment of the present invention.

[0019] The technical solution of this invention involves encapsulating an ultrasound transducer at the tip of an adhesive ultrasound probe. The acoustic window on the ultrasound transducer is planar and is used to attach to the surface of the heart or aorta. The plane of the ultrasound transducer is parallel to the plane of the acoustic window. The adhesive ultrasound probe can be stably attached to the surface of the heart structure in a specific posture, which solves the problem of acoustic interference from anatomical structures such as the chest wall and esophagus in traditional echocardiography. The direct coupling contact between the acoustic window and the surface of the heart structure significantly reduces the air gap, thereby improving the focusing accuracy and imaging efficiency of the ultrasound imaging system and providing real-time and high-quality image support for cardiac surgery.

[0020] It should be understood that the description in this section is not intended to identify key or essential features of the embodiments of the present invention, nor is it intended to limit the scope of the invention. Other features of the invention will become readily apparent from the following description. Attached Figure Description

[0021] To more clearly illustrate the technical solutions in the embodiments of the present invention, the accompanying drawings used in the description of the embodiments will be briefly introduced below. Obviously, the accompanying drawings described below are only some embodiments of the present invention. For those skilled in the art, other drawings can be obtained based on these drawings without creative effort.

[0022] Figure 1 This is a schematic diagram of the structure of an ultrasound imaging system based on the surface of the heart, provided in one embodiment of the present invention.

[0023] Figure 2 This is a schematic diagram of the structure of an ultrasonic detection component provided in one embodiment of the present invention;

[0024] Figure 3 This is a schematic diagram of the internal structure of an interventional sheath provided in one embodiment of the present invention;

[0025] Figure 4 This is a schematic diagram illustrating the combined use of an adhesive ultrasound probe, interventional sheath, and control handle according to an embodiment of the present invention.

[0026] Figure 5 This is a schematic diagram of the structure of a handheld part provided in one embodiment of the present invention;

[0027] Figure 6 A schematic diagram of a specific example of an ultrasound imaging system based on the surface of the heart provided in an embodiment of the present invention;

[0028] Figure 7 This is a schematic diagram illustrating the principle of a sound wave propagation path according to an embodiment of the present invention.

[0029] Figure 8 A flowchart illustrating an ultrasound imaging method based on the surface of the heart, as provided in one embodiment of the present invention;

[0030] Figure 9 A flowchart illustrating another method for ultrasound imaging based on the cardiac surface, provided as an embodiment of the present invention;

[0031] Figure 10 This is a schematic diagram of the fitted launch path under a first planning mode, a second planning mode, and a third planning mode, provided as an embodiment of the present invention. Detailed Implementation

[0032] To enable those skilled in the art to better understand the present invention, the technical solutions of the present invention will be clearly and completely described below with reference to the accompanying drawings of the embodiments of the present invention. Obviously, the described embodiments are only some embodiments of the present invention, and not all embodiments. Based on the embodiments of the present invention, all other embodiments obtained by those skilled in the art without creative effort should fall within the scope of protection of the present invention.

[0033] It should be noted that the terms "first," "second," etc., in the specification, claims, and accompanying drawings of this invention are used to distinguish similar objects and are not necessarily used to describe a specific order or sequence. It should be understood that such data can be interchanged where appropriate so that the embodiments of the invention described herein can be implemented in orders other than those illustrated or described herein. Furthermore, the terms "comprising" and "having," and any variations thereof, are intended to cover non-exclusive inclusion; for example, a process, method, system, product, or apparatus that comprises a series of steps or units is not necessarily limited to those steps or units explicitly listed, but may include other steps or units not explicitly listed or inherent to such processes, methods, products, or apparatus.

[0034] Figure 1 This is a schematic diagram of a cardiac surface-based ultrasound imaging system provided in one embodiment of the present invention. This embodiment can provide services for the cardiac surface-based ultrasound imaging method described in any embodiment of the present invention.

[0035] like Figure 1 As shown, the cardiac surface-based ultrasound imaging system includes: an ultrasound detection component 1 and an imaging terminal 2 connected in communication, wherein the ultrasound detection component 1 includes an attached ultrasound probe 10.

[0036] In this embodiment, the head end of the adhesive ultrasound probe 10 is encapsulated with an ultrasound transducer 11. The acoustic window 12 on the ultrasound transducer 11 is planar and is used to attach to the surface of the heart or aorta. The plane of the ultrasound transducer 11 is parallel to the plane of the acoustic window 12.

[0037] Specifically, the ultrasonic transducer 11 uses a piezoelectric crystal array as its core functional layer, with an impedance matching layer and an acoustic absorption backing layer stacked sequentially, and is entirely encapsulated at the head end of a sealed and protected adhesive ultrasonic probe 10. In an optional embodiment, the ultrasonic transducer 11 is a high-frequency broadband transducer or a phased array transducer. The ultra-high resolution of the high-frequency broadband transducer supports the generation of high-definition real-time images, meeting the needs for observing fine cardiac structures during cardiac surgery, and supporting various advanced ultrasound imaging modes such as harmonic imaging and pulse inversion imaging. The phased array transducer can achieve a wider field of view, eliminates the inertia and wear of mechanical movement, and has higher imaging efficiency.

[0038] In this embodiment, the planar acoustic window 12 allows the attached ultrasound probe 10 to be stably attached to the surface of the heart structure in a specific posture, solving the problem of acoustic interference from anatomical structures such as the chest wall and esophagus in traditional echocardiography. The direct coupling contact between the acoustic window 12 and the surface of the heart structure significantly reduces the air gap, thereby improving the focusing accuracy and imaging efficiency of the ultrasound imaging system. The attach-to-image detection structure eliminates the complex insertion process of the traditional TEE probe and avoids the trouble of repeatedly searching for the acoustic window in TTE technology. The operation is intuitive and fast, providing real-time and high-quality image support for cardiac surgery.

[0039] In an optional embodiment, the non-functional structure 13 of the attached ultrasonic probe 10, excluding the acoustic window 12, is a smooth curved surface. Exemplary examples include spheres, cylinders, and parabolic surfaces, but are not limited to the examples given above.

[0040] The advantage of this design is that it allows the attached ultrasound probe 10 to move flexibly within cavities or tissue spaces, thereby reducing the risk of damage to organs and tissues during imaging operations.

[0041] In the following embodiments of the present invention, "far end" and "near end" refer to the distance between the device and the imaging terminal 2. "Far end" is the end that is farther away from the imaging terminal 2, and "near end" is the end that is closer to the imaging terminal 2.

[0042] like Figure 2 As shown, in an optional embodiment, the ultrasound detection assembly 1 further includes an interventional sheath 20, which is connected to the tail end of the attached ultrasound probe 10, and the cavity of the interventional sheath 20 is used to accommodate the cable of the attached ultrasound probe 10.

[0043] The advantages of setting up the interventional sheath 20 are that, on the one hand, it enables a wired connection with the imaging terminal 2, thereby improving the stability of communication with the imaging terminal 2; on the other hand, it increases the detection depth of the attached ultrasound probe 10, thereby expanding the detection coverage of the attached ultrasound probe 10 on the cardiac structure.

[0044] like Figure 3 As shown, in an optional embodiment, the interventional sheath 20 is provided with an elastic support material 21 to achieve elastic adhesion between the adhesive ultrasound probe 10 and the surface of the heart or aorta.

[0045] During the imaging process, the attached ultrasound probe 10 can easily exert a certain amount of pressure on the heart, which may cause abnormalities in vital signs parameters, thereby increasing the operational risks of ultrasound imaging.

[0046] This embodiment, by employing an elastic interventional sheath 20, not only avoids excessive pressure on the heart when the attached ultrasound probe 10 is attached to the surface of the heart structure, reducing the operational risks of ultrasound imaging, but also allows the attached ultrasound probe 10 to dynamically follow the non-rigid movement of the heart when there is non-rigid movement, achieving real-time close contact with the surface of the heart structure, further ensuring the continuity and stability of ultrasound imaging.

[0047] like Figure 2 As shown, in an optional embodiment, the ultrasound detection assembly 1 further includes a control handle 30, the cable leading out of the interventional sheath 20 is connected to the imaging terminal 2 via the control handle 30, and the attached ultrasound probe 10 completes the position adjustment according to the control operation of the control handle 30.

[0048] The attached ultrasound probes used in TTE (Transthoracic Echocardiography) are large and limited by the size of the thoracic incision, restricting their flexibility within the chest cavity. They can only examine the heart from the front, creating blind spots and preventing the acquisition of images of the heart from other directions, such as the sides and back. TEE (Transthoracic Echocardiography) can only image from the back of the heart, also exhibiting certain imaging blind spots. For example, due to the distance between the esophagus and the right ventricle, TEE provides insufficient visualization of overall right ventricular function and outflow tract details, and also lacks depth in detecting the fine structures of the apex and anterior wall. Furthermore, TEE probes are more difficult to operate and require operators with advanced cardiac ultrasound qualifications.

[0049] In this embodiment, the control handle 30 can flexibly control the position and orientation of the attached ultrasound probe 10 within the chest cavity, reducing the difficulty of ultrasound imaging operation. It can be used in conjunction with other surgical equipment, and at the same time, it breaks through the directional limitations of traditional ultrasound imaging technology, expands the coverage of cardiac imaging, and thus ensures the integrity and accuracy of ultrasound imaging.

[0050] Figure 4 This is a schematic diagram illustrating the combined use of an adhesive ultrasound probe, an interventional sheath, and a control handle according to an embodiment of the present invention. The adhesive ultrasound probe 10, the interventional sheath 20, and the control handle 30 are coaxially connected. The adhesive ultrasound probe 10 achieves a flexible control link with the control handle 30 via the interventional sheath 20. The interventional sheath 20 is bent under the operating force from the control handle 30. The interventional sheath 20 transmits this operating force to the adhesive ultrasound probe 10, achieving a tight fit between the adhesive ultrasound probe 10 and the surface of the heart structure. The control operation of the control handle 30 can specifically be angle adjustment or displacement adjustment. The control operation of the control handle 30 is transmitted to the adhesive ultrasound probe 10 through the interventional sheath 20, driving the adhesive ultrasound probe 10 to dynamically adjust the attachment angle, attachment position, and attachment tightness on the surface of the heart structure, achieving a coordinated linkage effect of the control handle 30, the interventional sheath 20, and the adhesive ultrasound probe 10.

[0051] During the imaging procedure, the interventional sheath 20 can not only transmit the operating force applied by the control handle 30 to drive the attached ultrasound probe 10 to adjust its position, but also reduce the pressure component of the operating force on the heart, avoiding pressure on the heart caused by control errors of the control handle 30 or continuous pressing.

[0052] like Figure 2 As shown, in an optional embodiment, the control handle 30 includes a handheld part 31 and a connecting part 32. The distal end of the handheld part 31 is connected to the interventional sheath 20. The connecting part 32 includes an extension cable 321 and a tail plug 322. The proximal end of the handheld part 31 is connected to the imaging terminal 2 via the extension cable 321 and the tail plug 322 in sequence.

[0053] The extension cable 321 expands the operating range of the control handle 30, achieving spatial decoupling between the probe operating end and the imaging terminal 2. The control handle 30 allows the attached ultrasound probe 10 to move freely within the surgical field without concern for the placement of the imaging terminal 2. The tail plug 322 enables quick insertion and removal of the ultrasound detection component 1 and the imaging terminal 2, facilitating the replacement and sterilization of the ultrasound detection component 1.

[0054] like Figure 5 As shown, based on the above embodiment, optionally, the handheld part 31 is provided with a mark for identifying the plane orientation of the sound window 121. Figure 5 (not shown in the image), and / or, the handheld portion 31 includes a handheld structure 311 and a planar indicating structure 312, the plane of which is parallel to the plane of the sound window 121.

[0055] For example, the markings on the control handle 30 may be raised ridges, stickers with preset geometric patterns, or laser-etched arrow patterns, etc., wherein the preset geometric pattern may be a red triangle, and the color of the sticker must be different from the color of the control handle 30 itself.

[0056] like Figure 5 As shown, in an optional embodiment, the handheld structure 311 is further provided with a control button 33 and / or a groove 34 for accommodating fingers.

[0057] Specifically, groove 34 conforms to an ergonomic curve, the groove depth is adapted to the thickness of the fingers, and the inner wall of the groove can be made with an anti-slip texture.

[0058] For example, the control buttons 33 include, but are not limited to, pose adjustment buttons, imaging mode switching buttons, image capture buttons, parameter fine-tuning buttons, and emergency start / stop buttons, etc. Two or more control buttons 33 can be arranged linearly according to the operation frequency, such that the control buttons 33 closer to the groove 34 have a higher operation frequency.

[0059] The orientation of the acoustic window 121 is marked and indicated to avoid errors in the direction of the sound beam caused by the indirect connection between the attached ultrasound probe 10 and the control handle 30. The control button 33 enables convenient operation of functions such as position adjustment and imaging parameter setting of the attached ultrasound probe 10, improving the efficiency of cardiac surgery. The groove 34 improves the grip comfort of the control handle 30, thereby improving the stability during the imaging operation and further enhancing the safety of ultrasound imaging. The integrated application of the above hardware structures jointly ensures the efficient, accurate and stable operation of the ultrasound imaging system.

[0060] Based on the above embodiments, optionally, the adhesive ultrasound probe 10 is made of sterilizable material, and the ultrasound detection component 1 is designed for single use to avoid the risk of cross-infection.

[0061] In this embodiment, the imaging terminal 2 is used to generate a sound velocity mask based on a reference ultrasound image corresponding to the radio frequency signal data. The sound velocity mask contains at least two sound velocity distribution regions. The echo delay data is determined based on the sound velocity mask. Based on the echo delay data, the ultrasound image is synthesized and reconstructed from the radio frequency signal data. Motion compensation is performed on the image sequence composed of multiple ultrasound images. Multimodal fusion is performed on the motion-compensated ultrasound image to obtain the final ultrasound image.

[0062] like Figure 6 As shown, based on the above embodiments, optionally, the imaging terminal 2 includes a signal acquisition module 40, an image processing module 50, a data transmission module 60, a display module 70, and a power management module 80 for power supply.

[0063] The signal acquisition module 40 is used to acquire radio frequency signal data acquired by the attached ultrasound probe 10; the image processing module 50 is used to determine the ultrasound image corresponding to the radio frequency signal data; the data transmission module 60 is used to transmit the ultrasound image to the display module 70; and the display module 70 is used to output and display the ultrasound image on a visualization interface.

[0064] For example, the signal acquisition module 40 can also perform preprocessing such as filtering and amplification on radio frequency signal data; the image processing module 50 is an integrated high-performance processor that can complete image reconstruction, enhancement and dynamic adjustment in real time; the data transmission module 60 supports wired and wireless transmission methods and can transmit real-time image data to the display module 70; the display module 70 is equipped with a high-resolution display screen that supports real-time viewing of ultrasound images during surgery and provides image display in multiple ultrasound imaging modes; the power management module 80 can adopt a low-power design to ensure that the system can operate stably for a long time during cardiac surgery.

[0065] Based on the above embodiments, optionally, the imaging terminal 2 may also include a peripheral device 90 connected to the data transmission module 60. For example, the peripheral device 90 may include, but is not limited to, input units and output units. For example, input units may include keyboards, mice, microphones, cameras, handwriting tablets, touch screens, and card readers, and output units may include printers, projectors, headphones, speakers, and indicator lights, but are not limited to the given examples.

[0066] The tissue composition of the heart is complex and heterogeneous. For example, the acoustic properties of the myocardium, valves, blood, and even abnormal tissues are different, especially in terms of acoustic impedance (determined by both sound velocity and density). Figure 7 As shown, traditional image reconstruction techniques, based on the aforementioned assumptions, derive a straight-line sound wave propagation path. However, when ultrasound waves propagate within the heart structure, they pass through different biological tissues (such as…). Figure 7 The actual propagation speed will change when A, B, and C are in the heart. The propagation speed of sound waves is not uniformly distributed in the heart, and the actual propagation path of sound waves is tortuous.

[0067] As the imaging depth increases, the phase error introduced by the above approximation will gradually accumulate, causing distortions in ultrasound images such as geometric distortion and blurred tissue boundaries.

[0068] Figure 8 This is a flowchart illustrating a cardiac surface-based ultrasound imaging method according to one embodiment of the present invention. This embodiment is applicable to cardiac ultrasound imaging during cardiac surgery and can be used with the cardiac surface-based ultrasound imaging system described in any embodiment of the present invention.Figure 8 As shown, the method includes:

[0069] S210. Acquire radio frequency signal data collected by the attached ultrasound probe on the surface of the heart or aorta.

[0070] Among them, radio frequency signal data represents the parameter information of radio frequency electromagnetic waves received by the ultrasound transducer after ultrasound waves are reflected or scattered by biological tissue. The radio frequency signal data includes the radio frequency signal received by each receiving element in the ultrasound transducer, and the radio frequency signal carries the structural feature information of the heart.

[0071] S220. Generate a sound velocity mask based on the reference ultrasonic image corresponding to the radio frequency signal data.

[0072] In one alternative embodiment, the reference ultrasound image is acquired through at least one of the following methods:

[0073] 1) A historical frame ultrasound image reconstructed from radio frequency signal data corresponding to at least one historical acquisition time; a reference ultrasound image is determined based on at least one historical frame ultrasound image, wherein the historical acquisition time is located before the acquisition time of the radio frequency signal data.

[0074] 2) Using a constant sound wave propagation speed, a reference ultrasound image is obtained by synthesizing and reconstructing the radio frequency signal data.

[0075] Specifically, a time threshold is predefined for the time interval between the historical acquisition time and the acquisition time of the radio frequency signal data to ensure that there is a strong spatiotemporal correlation between the radio frequency signal data at the historical acquisition time and the radio frequency signal data, thereby ensuring the accurate constraint effect of the sound velocity information provided by the reference ultrasound image on the radio frequency signal data.

[0076] In one specific embodiment, the number of historical frame ultrasound images is one, and the historical frame ultrasound image is the previous frame ultrasound image corresponding to the radio frequency signal data at the previous acquisition time. Determining a reference ultrasound image based on at least one historical frame ultrasound image includes: using the previous frame ultrasound image as the reference ultrasound image.

[0077] In another specific embodiment, there are multiple historical frame ultrasound images. Determining a reference ultrasound image based on at least one historical frame ultrasound image includes: superimposing multiple historical frame ultrasound images according to image weight data to obtain the reference ultrasound image. The image weight data includes the image weights corresponding to each of the multiple historical frame ultrasound images. For example, the image weights corresponding to the multiple historical frame ultrasound images are the same, or the shorter the time interval between the historical acquisition time of the historical frame ultrasound image and the acquisition time of the radio frequency signal data to be reconstructed, the greater its corresponding image weight.

[0078] In this embodiment, the sound velocity mask contains at least two sound velocity distribution regions. The sound velocity mask represents the sound velocity information of the ultrasound at each imaging point corresponding to the reference ultrasound image. The imaging point has a one-to-one correspondence with the pixel in the reference ultrasound image. The imaging point represents the tissue target point within the detection range of the ultrasound transducer or in the heart.

[0079] In one optional embodiment, generating a sound velocity mask based on a reference ultrasound image corresponding to the radio frequency signal data includes: replacing the gray values ​​of imaging points in the reference ultrasound image with the sound wave propagation speed according to a sound velocity-grayscale mapping model to obtain the sound velocity mask. The sound velocity-grayscale mapping model is a pre-established model that maps gray values ​​to sound wave propagation speeds in a one-to-one correspondence. For example, the sound velocity-grayscale mapping model can be a linear mapping model, a logarithmic mapping model, or an adaptive nonlinear mapping model, but is not limited to the given examples.

[0080] In another optional embodiment, generating a sound velocity mask based on a reference ultrasound image corresponding to the radio frequency signal data includes: performing tissue segmentation on the reference ultrasound image to obtain a tissue segmentation image; and assigning values ​​to imaging points in the tissue segmentation area corresponding to the tissue type in the tissue segmentation image according to the sound wave propagation speed corresponding to the tissue type to obtain a sound velocity mask.

[0081] Specifically, the tissue segmentation image contains tissue segmentation regions corresponding to at least two tissue types. For example, the segmentation algorithm corresponding to the tissue segmentation operation is a threshold segmentation algorithm, a clustering segmentation algorithm, or a machine learning algorithm, but it is not limited to the given example.

[0082] Taking myocardium and blood as examples, the imaging points in the sound velocity mask image speed of sound wave propagation Satisfy the following formula:

[0083]

[0084] in, This indicates the speed at which sound waves travel in blood. This indicates the speed of sound wave propagation corresponding to the myocardium. Represents the imaging point Located within the blood segmentation area, Represents the imaging point Located within the myocardial segmentation region.

[0085] S230. Determine echo delay data based on the sound velocity mask, and synthesize and reconstruct an ultrasound image from the radio frequency signal data based on the echo delay data.

[0086] In this embodiment, the echo delay data includes the echo delay between each imaging point and each receiving array element. Specifically, for each imaging point and each receiving array element, the acoustic wave propagation path corresponding to the imaging point and the receiving array element is obtained, and the echo delay corresponding to the imaging point and the receiving array element is determined according to the acoustic wave propagation path and the sound velocity mask.

[0087] Here, the sound wave propagation path represents the complete path information of the ultrasonic wave from the transmitting element to the imaging point and the echo path from the imaging point back to the receiving element. For example, the imaging point... With the first in the ultrasonic transducer The echo delay corresponding to each receiving element Satisfy the following formula:

[0088]

[0089] in, This indicates the propagation step length of the ultrasound wave. This indicates the distance of the ultrasonic wave from the transmitting element to the imaging point. launch path, Indicates that the ultrasound waves originate from the imaging point. To the The echo path of each receiving array element Represents the imaging points in the sound speed mask image. The speed at which sound waves propagate.

[0090] Specifically, the sound wave propagation speed of the imaging point in the sound velocity mask is obtained, and the echo delays corresponding to the imaging point and multiple receiving array elements in the echo delay data are obtained. Based on the sound wave propagation speed and multiple echo delays, the imaging depth of the imaging point is determined. Combined with the array channel position and scanning angle of the ultrasonic transducer, the spatial coordinates of the imaging point are determined. Based on the multiple echo delays corresponding to the imaging point, the radio frequency signal sequence corresponding to the imaging point in the radio frequency signal data is weighted and superimposed to obtain the signal amplitude of the imaging point. Each signal amplitude is assigned to the pixel position of the corresponding spatial coordinate to generate an initial spatial amplitude matrix. Interpolation is performed on the gap region between adjacent imaging points in the initial spatial amplitude matrix to fill the pixel gaps and obtain the final spatial amplitude matrix. A grayscale mapping algorithm is used to map the final spatial amplitude matrix into a grayscale image of 0-255, i.e., an ultrasonic image.

[0091] For example, imaging points Corresponding signal amplitude Satisfy the following formula:

[0092]

[0093] in, Indicates the application of the first The window function weights on each receiving element are used to suppress sidelobe signals from other imaging points. Indicates the first The radio frequency signal received by each receiving array element This indicates the number of receiving array elements.

[0094] To achieve high frame rate and ultra-fast imaging, multi-beam parallel transmission or plane wave imaging technology can be used in the beamforming stage to improve imaging efficiency by reducing the number of transmission lines. For example, the reconstruction process can be accelerated by parallel processing on heterogeneous platforms of central processing unit and image processor using Compute Unified Device Architecture (CUDA), Open Computing Language (OpenCL), or OpenMP (Open Multi-Processing).

[0095] S240. Motion compensation is performed on the image sequence composed of multiple ultrasound images, and multimodal fusion is performed on the motion-compensated ultrasound images to obtain the final ultrasound image.

[0096] In high-speed ultrasound imaging, the heart undergoes non-rigid motion due to physiological activities such as breathing and heartbeat. At the same time, the attached ultrasound probe itself may also experience slight vibrations during operation. These movements can lead to problems such as image blurring and boundary artifacts in the image sequence, which is particularly noticeable in intraoperative scenarios where close contact with the tissue surface is required, thereby reducing the quality of ultrasound imaging.

[0097] Specifically, a motion estimation algorithm is used to determine the pixel displacement vector, tissue deformation law, or probe motion trajectory of the current frame ultrasound image in the image sequence compared to the previous frame ultrasound image. Based on the motion vector field, the imaging point coordinates of the imaging points in the current frame ultrasound image are compensated and corrected so that the tissue target points are aligned to the same positioning position in the image sequence.

[0098] For example, motion estimation algorithms include, but are not limited to, block matching motion estimation algorithms, optical flow field estimation algorithms, phase correlation methods, anatomical registration algorithms, or neural network algorithms, etc.

[0099] Taking the block matching motion estimation algorithm as an example, according to the image block size M×N, the current frame of ultrasound image is... The image is divided into current image patch sets, and the previous frame ultrasound image with motion compensation is also obtained. The previous image patch set is obtained by dividing the image into parts, and the search window size is defined as (M+2d)×(N+2d), where d represents the maximum possible displacement of the tissue target.

[0100] For the current image patch The algorithm retrieves multiple candidate image patches from the previous image patch set within the search window of the current image patch, and quantifies the similarity between the current image patch and the candidate image patches using a matching criterion. For example, the matching criterion value can be Mean Absolute Difference (MAD) or Mean Squared Error (MSE), but is not limited to the given example.

[0101] For example, the mean absolute error satisfies the following formula:

[0102]

[0103] The mean square error satisfies the following formula:

[0104]

[0105] in, This represents the previous ultrasound image after motion compensation. This represents the displacement vector, with a value range of 1000. .

[0106] Within the search region S, the matching criterion value between each candidate image block and the current image block is calculated. For example, when the matching criterion value is MAD or MSE, the candidate image block corresponding to the smallest MAD or MSE is taken as the matching image block similar to the current image block, and the difference between the center coordinates of the matching image block and the current image block is taken as the motion vector of the current image block.

[0107] Taking the matching criterion value of MAD as an example, the current image patch motion vector Satisfy the following formula:

[0108]

[0109] Motion vectors from multiple current image patches are integrated into a motion vector field, and motion compensation is performed on the current frame ultrasound image based on this motion vector field. For example, the motion-compensated current frame ultrasound image... Satisfy the following formula:

[0110]

[0111] in, This represents a moving vector field.

[0112] Ultrasound imaging mode refers to an imaging strategy based on the principles of sound wave propagation and signal processing. By adjusting the sound wave emission mode, beamforming algorithm, and signal analysis logic, it can achieve targeted imaging of different tissue anatomical structures and physiological functions to meet diverse imaging needs.

[0113] For example, ultrasound imaging modes include, but are not limited to, B-mode imaging (Brightness Mode), M-mode imaging (Motion Mode), Doppler ultrasound imaging, and elastography mode. B-mode imaging maps the acoustic impedance differences of biological tissues using grayscale brightness; for example, high echogenicity corresponds to dense tissues such as bone and calcifications, medium echogenicity corresponds to muscles and organ parenchyma, and low echogenicity corresponds to fluid tissues such as blood and effusion. M-mode imaging, based on B-mode imaging, selects a single ultrasound beam path, using time as the horizontal axis and depth as the vertical axis to record the dynamic trajectory of biological tissues, suitable for applications such as heart valve motion and ventricular wall systolic and diastolic function. Doppler ultrasound imaging utilizes the Doppler effect to quantify parameters such as blood flow velocity, direction, and volume by detecting frequency changes in the scattered echoes from red blood cells in the blood flow. Elastography detects the degree of tissue deformation by applying minute pressure to the tissue or using sound wave radiation, converting it into an elastic image, and achieving tissue identification based on differences in the elastic modulus of different tissues.

[0114] In this embodiment, the phase compensation process and motion compensation process provided in this embodiment are uniformly applied to multiple ultrasound imaging modes. The multiple ultrasound images in each image sequence are all compensated ultrasound images, so as to lay a precise spatial foundation for subsequent pixel-level fusion.

[0115] For example, multimodal fusion methods include, but are not limited to, weighted average fusion, multi-resolution pyramid fusion, wavelet transform-based image fusion, or deep learning models, etc.

[0116] Motion compensation in this embodiment suppresses the overall image jitter and blurring caused by heartbeat and respiration, further improving the quality of ultrasound imaging. At the same time, multimodal fusion can integrate complementary information from different imaging modes, ensuring that image data from different sources are highly consistent in time and space, reducing artifact interference from single-sequence imaging, and flexibly adapting to the imaging needs of different scenarios. Ultimately, it achieves accurate registration and seamless superposition of different information layers in anatomical position and dynamic features.

[0117] The technical solution of this embodiment distinguishes imaging points with sound velocity differences in the reference ultrasound image, constructs a sound velocity mask, combines sound wave path data to determine echo delay data, and synthesizes and reconstructs radio frequency signal data based on echo delay data to generate an ultrasound image. This solves the problem of echo delay deviating from the true value due to differences in acoustic characteristics, realizes phase compensation in the ultrasound imaging process, ensures the positioning accuracy of imaging points in the ultrasound image, and meets the application needs of high-precision imaging of cardiac structures in the medical field.

[0118] Figure 9This is a flowchart of another ultrasound imaging method based on the cardiac surface provided in one embodiment of the present invention. This embodiment further refines the "determining echo delay data based on the sound velocity mask" in the above embodiment. In this embodiment, determining the echo delay data based on the sound velocity mask includes: for each imaging point in the sound velocity mask, determining the sound wave propagation path corresponding to each imaging point and each receiving array element by combining a path planning mode; and determining the echo delay data based on multiple sound wave propagation paths and the sound velocity mask. Figure 9 As shown, the method includes:

[0119] S310. Acquire radio frequency signal data collected by the attached ultrasound probe on the surface of the heart or aorta.

[0120] S310 in this embodiment is the same as that in the above embodiment. Figure 8 The S210 shown is similar to or corresponds to the S210 shown, and will not be described again in this embodiment.

[0121] In an optional embodiment, the method further includes: acquiring multiple frames of radio frequency signals repeatedly collected at the blood flow detection point, and acquiring the sound wave propagation speed at the blood flow detection point from the sound speed mask; determining the inter-frame phase difference based on the multiple frames of radio frequency signals; and determining the blood flow velocity at the blood flow detection point based on the inter-frame phase difference and the sound wave propagation speed.

[0122] Specifically, inter-frame differentiation is performed on multiple frames of radio frequency signals to obtain signal change rate data. After high-pass filtering of the signal change rate data, the filtered signal change rate data is orthogonally demodulated to obtain orthogonal baseband IQ (In-phase-Quadrature) signal data. Secondary wall filtering is performed on the orthogonal baseband IQ signal data to filter out blood vessel wall interference signals. Based on the orthogonal baseband IQ signal data obtained after filtering, the inter-frame phase difference is determined.

[0123] For example, inter-frame phase difference Satisfy the following formula:

[0124]

[0125] in, Let be the single-hysteresis autocorrelation function, representing the average value of the conjugate product of each orthogonal baseband IQ signal and its next orthogonal baseband IQ signal. Indicates the number of radio frequency signals collected. Indicates the first The corresponding orthogonal baseband IQ signal was acquired in the next acquisition. Indicates the first The corresponding orthogonal baseband IQ signal was acquired in the next acquisition.

[0126] For example, the blood flow velocity at the location q to be detected Satisfy the following formula:

[0127]

[0128] in, Indicates the transmission frequency.

[0129] In color Doppler or pulsed Doppler modes, the heart's movement introduces interference with the velocity spectrum. This embodiment uses a notch filter and a narrowband high-pass filter to filter out specific frequency or low-frequency signals, thereby further improving the accuracy of blood flow velocity.

[0130] S320. Generate a sound velocity mask based on the reference ultrasonic image corresponding to the radio frequency signal data.

[0131] S320 in this embodiment is the same as that in the above embodiment. Figure 8 The S220 shown is similar to or corresponds to the S220 shown, and will not be described again in this embodiment.

[0132] S330. For each imaging point in the sound velocity mask, the sound wave propagation path corresponding to the imaging point and each receiving array element is determined by combining the path planning mode.

[0133] In this embodiment, the path planning rates corresponding to the first planning mode, the second planning mode, and the third planning mode in the path planning mode increase sequentially.

[0134] In an optional embodiment, the acoustic propagation path corresponding to the imaging point and each receiving array element is determined by combining a path planning mode, including: for each receiving array element, obtaining the coordinates of the receiving array element, the coordinates of the transmitting array element matching the receiving array element, and the coordinates of the imaging point; in response to the path planning mode being a first planning mode, using a ray tracing algorithm, determining an initial propagation path based on the coordinates of the receiving array element, the coordinates of the transmitting array element, and the coordinates of the imaging point, and iteratively optimizing the initial propagation path according to iterative constraints to obtain the acoustic propagation path corresponding to the imaging point and the receiving array element; in response to the... The path planning mode is the second planning mode, which divides the sound velocity mask into a grid to obtain a discrete sound velocity map. The shortest propagation time is used as a constraint. Path search is performed based on the coordinates of the receiving array element, the transmitting array element, and the imaging point to obtain the sound wave propagation path corresponding to the imaging point and the receiving array element. In response to the path planning mode being the third planning mode, the intersection sequence corresponding to at least two sound velocity distribution regions is determined based on the coordinates of the receiving array element, the transmitting array element, and the imaging point. Based on the intersection sequence, the sound wave propagation path corresponding to the imaging point and the receiving array element is determined.

[0135] For example, the triggering conditions corresponding to the first planning mode include, but are not limited to, receiving an operation instruction corresponding to the first planning mode, the system's computing power load being less than the minimum load threshold, the current moment satisfying the triggering period corresponding to the first planning mode, and the data quality corresponding to the radio frequency signal data being higher than the maximum data quality threshold, but are not limited to the example situations given above.

[0136] In this embodiment, the initial propagation path represents the straight path of the ultrasonic wave starting from the transmitting element, passing through the imaging point, and finally reaching the receiving element. The path planning accuracy of the first planning mode is the pixel size. The iterative constraints include the sound velocity mask and the node spacing in the sound wave propagation path not being greater than the pixel size. The sound velocity mask is used to constrain the sound velocity propagation path to satisfy the law of refraction of ultrasonic waves along the direction of the sound velocity gradient.

[0137] For example, the iterative optimization process may employ iterative algorithms including, but not limited to, Newton's iteration method or the steepest descent method. For instance, based on the node spacing in the iterative constraints, the initial propagation path is discretized into multiple path nodes. According to the sound velocity mask, the constraint deviation value for each path node is determined. The sound wave propagation path is iteratively obtained by minimizing the constraint deviation value as the objective function.

[0138] For example, the triggering conditions corresponding to the second planning mode include, but are not limited to, receiving an operation instruction corresponding to the second planning mode, the system's computing power load being greater than or equal to the minimum load threshold and less than the maximum load threshold, the current moment satisfying the triggering period corresponding to the second planning mode, and the data quality corresponding to the radio frequency signal data being lower than or equal to the maximum data quality threshold and higher than the minimum data quality threshold, etc., but are not limited to the example situations given above.

[0139] In this embodiment, the path planning accuracy of the second planning mode is the grid size. For example, the search algorithm used in the path search process includes, but is not limited to, fast traversal method, single-source shortest path algorithm or heuristic shortest path search algorithm, etc.

[0140] For example, the triggering conditions corresponding to the third planning mode include, but are not limited to, receiving an operation instruction corresponding to the third planning mode, the system's computing power load being greater than the maximum load threshold, the current moment satisfying the triggering period corresponding to the third planning mode, and the data quality corresponding to the radio frequency signal data being lower than the minimum quality threshold, but are not limited to the example situations given above.

[0141] In this embodiment, the path planning accuracy of the third planning mode is the path length of the biological tissue through which the ultrasonic wave passes along the propagation direction.

[0142] Specifically, the intersection sequence includes the transmission intersection sequence and the echo intersection sequence. The transmission intersection sequence represents the intersection points generated by the ultrasonic wave as it passes through different sound velocity distribution areas in the sound velocity mask after being emitted from the transmitting array element. These intersection points are arranged in order of the ultrasonic wave's emission direction to form an ordered coordinate set. The echo intersection sequence represents the intersection points generated by the ultrasonic wave as it passes through different sound velocity distribution areas in the sound velocity mask from the imaging point to the receiving array element. These intersection points are arranged in order of the ultrasonic wave's echo direction to form an ordered coordinate set.

[0143] Figure 10 This is a schematic diagram of the fitted transmission path under a first planning mode, a second planning mode, and a third planning mode provided in an embodiment of the present invention. Specifically, the more segments on the transmission path, the higher the path planning accuracy of the transmission path, and correspondingly, the lower the path planning rate of the sound wave propagation path. Conversely, the fewer segments on the transmission path, the lower the path planning accuracy of the transmission path, and correspondingly, the higher the path planning efficiency of the sound wave propagation path.

[0144] Specifically, the first and second planning modes can significantly reduce focusing errors caused by the non-uniform distribution of sound velocity. The third planning mode sacrifices the spatial resolution of ultrasound images within an acceptable range in exchange for higher temporal resolution of ultrasound images to meet the rapid imaging requirements of cardiac surgery.

[0145] In another optional embodiment, determining the acoustic wave propagation path corresponding to the imaging point and each receiving array element, in conjunction with the path planning mode, includes: in response to the path planning mode being the third planning mode, for each receiving array element, searching for an ultrasound image template that is similar to and matches the reference ultrasound image from the ultrasound image template set, obtaining a path lookup table corresponding to the ultrasound image template, the path lookup table containing the acoustic wave propagation path of each receiving array element in the ultrasound transducer under each imaging anchor point corresponding to the ultrasound image template; and determining the acoustic wave propagation path corresponding to the imaging point and the receiving array element based on the imaging point coordinates of the imaging point and the path lookup table.

[0146] For example, the ultrasound image template set can be associated with specific anatomical structures or specific imaging scenes, or with specific ultrasound imaging parameters such as probe model, imaging depth, and imaging angle. Imaging anchor points are imaging points that are pre-selected in the ultrasound image template for positioning, calibration, registration, or tracking. They have characteristics such as high recognition, spatial stability, path representativeness, and repeatability. Their purpose is to provide a reference benchmark for path calculation for a large number of other imaging points in the same template scene. The set of imaging anchor points constitutes the core index of the path lookup table.

[0147] Specifically, the matched ultrasound image template is the template image with the highest image similarity between the ultrasound image template set and the reference ultrasound image. For example, similarity matching algorithms include, but are not limited to, gray-level histogram matching algorithms, scale-invariant feature transformation algorithms, or convolutional neural network algorithms, etc.

[0148] The advantage of setting up similarity matching and path lookup tables is that by simplifying the real-time calculation steps of sound wave propagation paths, the efficiency of path planning is greatly improved, and the problem of computing power overload caused by full-pixel path calculation is effectively avoided. It can adapt to different path planning needs and reduce the hardware resource consumption of real-time computing, thus achieving the optimal balance between imaging speed and image quality in all application scenarios.

[0149] In an optional embodiment, determining the acoustic wave propagation path corresponding to the imaging point and the receiving array element based on the imaging point coordinates and the path lookup table includes: retrieving an imaging anchor point with the same coordinates as the imaging point in the path lookup table, and using the acoustic wave propagation path corresponding to the imaging anchor point and the receiving array element as the acoustic wave propagation path corresponding to the imaging point and the receiving array element.

[0150] In another optional embodiment, determining the acoustic wave propagation path corresponding to the imaging point and the receiving array element based on the imaging point coordinates and the path lookup table includes: if the path lookup table does not contain the imaging point, searching the path lookup table for the imaging anchor point closest to the imaging point coordinates, and using the acoustic wave propagation path corresponding to the imaging anchor point and the receiving array element as the acoustic wave propagation path corresponding to the imaging point and the receiving array element.

[0151] In another optional embodiment, determining the acoustic wave propagation path corresponding to the imaging point and the receiving array element based on the imaging point coordinates and the path lookup table includes: if the imaging point is not included in the path lookup table, determining a first imaging anchor point with an imaging depth less than the imaging point and a second imaging anchor point with an imaging depth greater than the imaging point in the path lookup table based on the imaging point coordinates; searching for a first propagation path of the array element under the first imaging anchor point and a second propagation path of the array element under the second imaging anchor point from the path lookup table; and performing linear interpolation based on the first propagation path and the second propagation path to obtain the acoustic wave propagation path corresponding to the imaging point and the receiving array element.

[0152] For example, the number of imaging anchors can be 256, and different imaging anchors correspond to different imaging depths. In the segmented interval between two imaging anchors separated by N imaging points, the distance between adjacent imaging points and the first... The delay difference corresponding to each receiving element To satisfy the following formula:

[0153]

[0154] in, Indicates that one of the imaging anchor points is related to the first The echo delay corresponding to each receiving array element Indicates another imaging anchor point and the first The echo delay corresponding to each receiving array element.

[0155] The advantage of building a small number of imaging anchors is that it can significantly reduce the real-time computational load and achieve a balance between imaging speed and image quality with lightweight storage costs.

[0156] S340. Based on the multiple sound wave propagation paths and the sound velocity mask, determine the echo delay data, and based on the echo delay data, synthesize and reconstruct the radio frequency signal data to obtain an ultrasonic image.

[0157] S350. Motion compensation is performed on the image sequence composed of multiple ultrasound images, and multimodal fusion is performed on the motion-compensated ultrasound images to obtain the final ultrasound image.

[0158] S340-S350 in this embodiment are the same as those in the above embodiment. Figure 8 The S230-S240 shown are corresponding to or similar to those shown, and will not be described again in this embodiment.

[0159] The technical solution of this embodiment, by pre-establishing multiple path planning modes with varying efficiencies and combining these modes, determines the acoustic wave propagation path between the imaging point and the receiving array element. This solves the problem of the acoustic wave propagation path deviating from the actual path due to idealized assumptions. It can achieve high-precision path planning with limited computing power, reducing hardware dependence on computing power and storage. It allows for on-demand switching between lightweight and efficient or fine and precise path calculation logic, avoiding the waste of computing power or positioning deviations in a single planning mode, and balancing the dual technical requirements of real-time performance and accuracy in ultrasound imaging.

[0160] Based on the above embodiments, optionally, the method further includes: obtaining the image quality score corresponding to the final ultrasound image; and in response to the image quality score being less than a standard score threshold, outputting information for prompting adjustment of the attachment pressure and / or attachment angle of the adhesive ultrasound probe until the image quality score is greater than or equal to the standard score threshold.

[0161] The image quality score represents the result of quantitative evaluation of the ultrasound image. For example, the image quality score can be determined by methods such as the gray-level co-occurrence matrix method based on statistical features, the Sobel operator method based on edge features, the statistical modeling method based on signal-to-noise ratio, the image quality evaluation model, external equipment indicators, structural similarity with the standard ultrasound template, peak signal-to-noise ratio, etc., but is not limited to the examples given above.

[0162] The standard score threshold represents the quality pass line set to meet imaging requirements. For example, the information output methods include, but are not limited to, text, icons, or voice.

[0163] Specifically, insufficient adhesion pressure can create an air gap between the ultrasound probe and the heart or aorta, preventing effective penetration of ultrasound waves and leading to strong reflection interference. Excessive adhesion pressure may compress the heart, causing deformation such as narrowing of blood vessels, thus affecting the accuracy of the ultrasound image. An incorrect adhesion angle may prevent ultrasound waves from entering the heart perpendicularly, resulting in refraction and scattering phenomena that affect the clarity of the ultrasound image.

[0164] The benefit of setting quality prompts is that it reduces the error of human judgment and the operational threshold, especially in primary healthcare or navigation scenarios, by guiding operational behavior to achieve quality control of ultrasound imaging.

[0165] Various embodiments of the systems and techniques described above herein can be implemented in the following systems or combinations thereof: digital electronic circuit systems, integrated circuit systems, field-programmable gate arrays (FPGAs), application-specific integrated circuits (ASICs), application-specific standard parts (ASSPs), system-on-chips (SOCs), complex programmable logic devices (CPLDs), computer hardware, firmware, software, and / or combinations thereof. These various embodiments may include implementations in one or more computer programs that can be executed and / or interpreted on a programmable system including at least one programmable processor, which may be a dedicated or general-purpose programmable processor, capable of receiving data and instructions from a storage system, at least one input device, and at least one output device, and transmitting data and instructions to the storage system, the at least one input device, and the at least one output device.

[0166] Computer programs for implementing the cardiac surface-based ultrasound imaging method of the present invention can be written in any combination of one or more programming languages. These computer programs can be provided to a processor of a general-purpose computer, a special-purpose computer, or other programmable data processing device, such that when executed by the processor, the computer programs cause the functions / operations specified in the flowcharts and / or block diagrams to be implemented. The computer programs can be executed entirely on a machine, partially on a machine, as a standalone software package partially on a machine and partially on a remote machine, or entirely on a remote machine or server.

[0167] In the context of this application, a computer-readable storage medium can be a tangible medium that may contain or store a computer program for use by or in conjunction with an instruction execution system, apparatus, or device. A computer-readable storage medium can be, but is not limited to, electronic, magnetic, optical, electromagnetic, infrared, or semiconductor systems, apparatus, or devices, or any suitable combination of the foregoing. Alternatively, a computer-readable storage medium can be a machine-readable storage medium. Examples of machine-readable storage media include, based on an electrical connection of at least one wire, a portable computer disk, a hard disk, random access memory (RAM), read-only memory (ROM), erasable programmable read-only memory (EPROM), flash memory, optical fiber, compact disc read-only memory (CD-ROM), optical storage devices, magnetic storage devices, or any suitable combination of the foregoing.

[0168] To provide user interaction, the systems and techniques described herein can be implemented on an imaging terminal having: a display device for displaying information to the user (e.g., a cathode-ray tube (CRT) or liquid crystal display (LCD) monitor); and a keyboard and pointing device (e.g., a mouse or trackball) through which the user provides input to the imaging terminal. Other types of devices can also provide user interaction; for example, feedback provided to the user can be any form of sensory feedback (e.g., visual feedback, auditory feedback, or tactile feedback); and input from the user can be received in any form (including sound input, voice input, or tactile input).

[0169] The systems and technologies described herein can be implemented in computing systems that include backend components (e.g., as data servers), or middleware components (e.g., application servers), or frontend components (e.g., user computers with graphical user interfaces or web browsers through which users can interact with implementations of the systems and technologies described herein), or any combination of such backend, middleware, or frontend components. The components of the system can be interconnected via digital data communication of any form or medium (e.g., communication networks). Examples of communication networks include local area networks (LANs), wide area networks (WANs), blockchain networks, and the Internet.

[0170] A computing system can include clients and servers. Clients and servers are generally located far apart and typically interact through communication networks. The client-server relationship is created by computer programs running on the respective computers and having a client-server relationship with each other. The server can be a cloud server, also known as a cloud computing server or cloud host, which is a hosting product within the cloud computing service system. It addresses the shortcomings of traditional physical hosts and Virtual Private Server (VPS) services, such as high management difficulty and weak business scalability.

[0171] It should be understood that the various forms of processes shown above can be used, with steps reordered, added, or deleted. For example, the steps described in this invention can be executed in parallel, sequentially, or in different orders, as long as the desired result of the technical solution of this invention can be achieved, and no limitation is imposed herein.

[0172] The specific embodiments described above do not constitute a limitation on the scope of protection of this invention. Those skilled in the art should understand that various modifications, combinations, sub-combinations, and substitutions can be made according to design requirements and other factors. Any modifications, equivalent substitutions, and improvements made within the spirit and principles of this invention should be included within the scope of protection of this invention.

Claims

1. An ultrasound imaging system based on a heart surface, characterized by, The application relates to a heart-surface-based ultrasonic imaging system. The head end of the attached ultrasonic probe is packaged with an ultrasonic transducer, and the sound window on the ultrasonic transducer is flat and used for being attached to the surface of a heart or an aorta; the plane of the ultrasonic transducer is parallel to the plane of the sound window. The imaging terminal is used for receiving the radio frequency signal data collected by the attached ultrasonic probe, generating a sound velocity mask image according to a reference ultrasonic image corresponding to the radio frequency signal data, containing at least two sound velocity distribution regions in the sound velocity mask image, determining echo time delay data according to the sound velocity mask image, synthesizing and reconstructing an ultrasonic image according to the echo time delay data, performing motion compensation on an image sequence formed by multiple ultrasonic images, performing multi-modal fusion on the motion-compensated ultrasonic image, and obtaining a final ultrasonic image. The non-functional structure of the attached ultrasonic probe except the sound window is a smooth curved surface.

2. The heart surface based ultrasound imaging system of claim 1, wherein, The ultrasonic detection assembly further comprises an interventional sheath connected with the tail end of the attached ultrasonic probe, and the cavity of the interventional sheath is used for accommodating the cable of the attached ultrasonic probe.

3. The heart surface based ultrasound imaging system of claim 1, wherein, The interventional sheath is provided with an elastic supporting material, which is used for realizing the elastic attachment of the attached ultrasonic probe to the surface of the heart or the aorta.

4. The heart surface based ultrasound imaging system of claim 3, wherein, The ultrasonic detection assembly further comprises a control handle, the cable led out of the interventional sheath is connected with the imaging terminal through the control handle, and the attached ultrasonic probe completes pose adjustment through the control operation of the control handle.

5. The heart surface based ultrasound imaging system of claim 3, wherein, The control handle comprises a hand-held part and a connecting part, the distal end of the hand-held part is connected with the interventional sheath, the connecting part comprises an extension cable and a tail plug, and the proximal end of the hand-held part is connected to the imaging terminal through the extension cable and the tail plug in sequence.

6. The heart surface based ultrasound imaging system of claim 5, wherein, The hand-held part is provided with a mark for identifying the plane orientation of the sound window, and / or the hand-held part further comprises a hand-held structure and a flat indicating structure, and the plane of the indicating structure is parallel to the plane of the sound window.

7. The heart surface based ultrasound imaging system of claim 6, wherein, The hand-held structure is further provided with a control button and / or a groove for accommodating fingers.

8. The heart surface based ultrasound imaging system of claim 7, wherein, The imaging terminal comprises a signal acquisition module, an image processing module, a data transmission module, a display module and a power management module for power supply.

9. The ultrasound imaging system based on the surface of the heart according to any one of claims 1 to 8, characterized in that, The signal acquisition module is used for acquiring the radio frequency signal data collected by the attached ultrasonic probe, the image processing module is used for determining an ultrasonic image corresponding to the radio frequency signal data, the data transmission module is used for transmitting the ultrasonic image to the display module, and the display module is used for outputting and displaying the ultrasonic image on a visual interface. The application is applied to the heart-surface-based ultrasonic imaging system in any one of claims 1-9.

10. An ultrasound imaging method based on a heart surface, characterized in that, The radio frequency signal data collected by the attached ultrasonic probe on the surface of a heart or an aorta is acquired. A sound velocity mask image is generated according to a reference ultrasonic image corresponding to the radio frequency signal data, and the sound velocity mask image contains at least two sound velocity distribution regions. ​ determine echo time delay data according to the sound velocity mask map, and synthesize and reconstruct the radio frequency signal data to obtain an ultrasound image according to the echo time delay data; perform motion compensation on an image sequence composed of multiple ultrasound images, perform multi-modal fusion on the motion-compensated ultrasound images, and obtain a final ultrasound image.

11. The heart surface based ultrasound imaging method of claim 10, wherein, The method further comprises: For each imaging point in the sound velocity mask map, a sound wave propagation path corresponding to each receiving array element is determined in combination with a path planning mode; According to the multiple sound wave propagation paths and the sound velocity mask map, echo time delay data is determined. The path planning rates of the first planning mode, the second planning mode, and the third planning mode in the path planning mode are sequentially increased.

12. The heart surface based ultrasound imaging method of claim 11, wherein, The method further comprises: For each receiving array element, the coordinates of the receiving array element, the coordinates of a transmitting array element matched with the receiving array element, and the coordinates of the imaging point are obtained; In response to the path planning mode being the first planning mode, a ray tracing algorithm is used to determine an initial propagation path according to the coordinates of the receiving array element, the coordinates of the transmitting array element, and the coordinates of the imaging point, and the initial propagation path is iteratively optimized according to an iterative constraint condition to obtain the sound wave propagation path corresponding to the receiving array element and the imaging point; In response to the path planning mode being the second planning mode, the sound velocity mask map is grid-divided to obtain a discrete sound velocity map, a shortest propagation time is used as a constraint condition, and path searching is performed according to the coordinates of the receiving array element, the coordinates of the transmitting array element, and the coordinates of the imaging point to obtain the sound wave propagation path corresponding to the receiving array element and the imaging point; In response to the path planning mode being the third planning mode, a sequence of intersection points corresponding to at least two sound velocity distribution regions is determined according to the coordinates of the receiving array element, the coordinates of the transmitting array element, and the coordinates of the imaging point, and the sound wave propagation path corresponding to the receiving array element and the imaging point is determined according to the sequence of intersection points.

13. The heart surface based ultrasound imaging method of claim 10, wherein, The method further comprises: Multiple frames of radio frequency signals repeatedly collected at a blood flow detection point are obtained, and a sound wave propagation velocity at the blood flow detection point is obtained from the sound velocity mask map; An inter-frame phase difference is determined according to the multiple frames of radio frequency signals; A blood flow velocity at the blood flow detection point is determined according to the inter-frame phase difference and the sound wave propagation velocity.

14. The heart surface based ultrasound imaging method of claim 10, wherein, The method further comprises: An image quality score corresponding to the final ultrasound image is obtained; In response to the image quality score being less than a standard score threshold, information for prompting adjustment of an attachment pressure and / or an attachment angle of the attached ultrasound probe is output until the image quality score is greater than or equal to the standard score threshold.

15. A computer-readable storage medium, characterized in that, The computer-readable storage medium stores computer instructions for causing a processor to implement the method of ultrasound imaging based on a heart surface according to any one of claims 10-14 when executed. The computer-readable storage medium stores computer instructions for causing a processor to implement the method of ultrasound imaging based on a heart surface according to any one of claims 10-14 when executed.

16. A computer program product comprising a computer program which, when executed by a processor, implements the heart surface based ultrasound imaging method according to any one of claims 10-14.

Images