Imaging die body assembly and image generation method

Abstract

One or more embodiments of the present specification relate to an imaging motif assembly and an image generation method, the imaging motif assembly comprising: a motif body; and at least one part of at least one coil of the one or more coils is arranged in the die body main body.

Description

Technical Field

[0001] The present invention relates to the field of imaging correction technology, and in particular to an imaging phantom component and an image generation method. Background Art

[0002] In the process of research, development, design and production of medical equipment (for example, MRI equipment), due to the limitations of physical design and manufacturing process, the main magnetic field generated will be integrated with the inhomogeneity of the magnetic field and the nonlinearity of the gradient generated in the gradient field. This influence will further introduce geometric distortion into the imaging space of the MRI equipment, and then cause geometric distortion of the imaging image. Therefore, it is necessary to use a phantom to correct the geometric distortion of the imaging space. In the process of correcting the distortion characteristics of large-scale imaging space, due to the large volume of the phantom required for geometric distortion, the way the RF receiving coil covers the phantom limits the signal quality of the central area of ​​the phantom that the RF coil can obtain, and it is impossible to achieve high-resolution, high-signal-to-noise ratio geometric distortion correction imaging.

[0003] In order to solve the above problems, the present specification provides an imaging phantom assembly and an image generation method, which can ensure that the signal in the central area of ​​the phantom body acquired by the RF coil also has a high quality, so as to achieve high-resolution, high signal-to-noise ratio geometric distortion correction imaging, further improve the geometric distortion correction effect, and ultimately improve the imaging quality of medical equipment. Summary of the invention

[0004] One of the embodiments of the present specification provides an imaging phantom assembly, including: a phantom body; and one or more coils, wherein at least a portion of at least one coil of the one or more coils is disposed in the phantom body.

[0005] One of the embodiments of the present specification provides an image generation method, comprising: acquiring a magnetic resonance signal of the imaging phantom assembly as described in the above embodiment under a radio frequency signal; and generating an image based on the magnetic resonance signal.

[0006] The imaging phantom assembly provided by the present invention can be applied to the geometric distortion correction of spatial uniform field in the medical and physical fields, so that the signals of various parts of the imaging phantom acquired by the radio frequency coil have high quality and accuracy, thereby obtaining high-quality and high-accuracy magnetic resonance signals, so that the imaging effect of the corrected image finally generated for geometric distortion correction is no longer limited by the coating method of the radio frequency coil on the phantom, thereby achieving the purpose of high-resolution and high signal-to-noise ratio geometric distortion correction imaging, and further improving the geometric distortion correction effect. BRIEF DESCRIPTION OF THE DRAWINGS

[0007] This specification will be further described in the form of exemplary embodiments, which will be described in detail by the accompanying drawings. These embodiments are not restrictive, and in these embodiments, the same number represents the same structure, wherein:

[0008] Figure 1 is a simplified module schematic diagram of an imaging phantom assembly according to some embodiments of this specification;

[0009] Figure 2 is a schematic diagram of an imaging phantom assembly according to some embodiments of the present specification;

[0010] Figure 3 is a schematic diagram of a single-layer structure of an imaging phantom assembly according to some embodiments of this specification;

[0011] Figure 4 is a schematic structural diagram of an imaging phantom assembly according to other embodiments of the present specification;

[0012] Figure 5 is a schematic structural diagram of a second model layer according to other embodiments of this specification;

[0013] Figure 6 is an exemplary flow chart of an image generating method according to some other embodiments of the present specification. DETAILED DESCRIPTION

[0014] In order to more clearly illustrate the technical solutions of the embodiments of this specification, the following is a brief introduction to the drawings required for the description of the embodiments. Obviously, the drawings described below are only some examples or embodiments of this specification. For ordinary technicians in this field, this specification can also be applied to other similar scenarios based on these drawings without creative work. Unless it is obvious from the language environment or otherwise explained, the same reference numerals in the figures represent the same structure or operation.

[0015] As shown in this specification and claims, unless the context clearly indicates an exception, the words "a", "an", "an" and / or "the" do not refer to the singular and may also include the plural. Generally speaking, the terms "comprise" and "include" only indicate the inclusion of the steps and elements that have been clearly identified, and these steps and elements do not constitute an exclusive list. The method or device may also include other steps or elements.

[0016] In the process of developing, designing and producing MRI equipment (i.e. magnetic resonance imaging equipment), due to the limitations of physical design and manufacturing process, the inhomogeneity of magnetic field and nonlinearity of gradient are usually incorporated into the generated main magnetic field and gradient field. This effect will introduce geometric distortion in the imaging space of MRI, which will lead to geometric distortion in the image, that is, the deviation of the coordinate points displayed in the MRI image from the known position or the incorrect ratio of the distance between any two points in the MRI image.

[0017] In some schemes, the technology for dealing with field inhomogeneity usually includes collecting magnetic field intensity at each point in the spatial shim (e.g., the magnetic field space corresponding to the imaging space) through a magnetic field intensity measurement tool, and correcting the uniformity of the magnetic field through ferromagnetic materials such as silicon steel sheets based on the spatial distribution of the magnetic field intensity. However, the method of shim correction is relatively rough and it is difficult to meet the geometric distortion correction requirements of high-resolution and high-precision MRI imaging such as radiotherapy. In addition, the scheme can only roughly compensate for the uniformity of the spatial shim surface (e.g., the spherical shim surface) during actual operation. The accuracy of the magnetic field intensity at each point in the spatial shim actually collected is low, and the compensation accuracy cannot meet the requirements. Therefore, in the case of high requirements for imaging resolution and geometric accuracy, it is necessary to place a phantom (the surface of the phantom is covered with a radio frequency coil) for geometric distortion correction in the spatial shim, and image the phantom to obtain a corrected image, and then present the distortion characteristics in the spatial shim based on the geometric structure in the corrected image, and then correct the geometric distortion. In some schemes, when imaging a large-sized phantom, the types of coils used to cover the phantom mainly include VTC volume transmit coils (Volume Transmit Coil) or BAC body array coils (Body Array Coil), SPC spine coils (Spine Coil), etc. or their combination. However, the actual imaging effect of the corrected image obtained is subject to the way the RF coil covers the phantom. In the actual sampling process, the signal quality of the central area of ​​the phantom obtained by the RF receiving coil is poor, which is not convenient for the identification and calculation of geometric distortion. In order to solve the limitation of the way the RF coil covers the phantom on the actual imaging effect, some schemes will use small-sized phantoms for imaging. However, the size of the spatial shim that can be corrected by this method is limited by the size of the phantom volume. For example, when the size of the spatial shim to be corrected is large, the small-sized phantom cannot cover the entire spatial shim, and thus cannot achieve geometric distortion correction. If the phantom is moved to compensate, it may bring more mechanical tail swing and fusion errors between the corrected image.

[0018] The imaging phantom assembly provided in the embodiments of this specification can be applied to the geometric distortion correction of the spatial uniform field in the medical and physical fields. Through the imaging phantom assembly provided in this specification, the signals of each part of the imaging phantom acquired by the radio frequency coil can have high quality and accuracy, thereby obtaining high-quality and high-accuracy magnetic resonance signals, so that the imaging effect of the correction image finally generated for geometric distortion correction is no longer limited by the coating method of the radio frequency coil on the phantom, thereby achieving the purpose of high-resolution and high signal-to-noise ratio geometric distortion correction imaging, and further improving the geometric distortion correction effect. In a specific application scenario, the imaging phantom assembly can be applied to the geometric distortion correction of MRI equipment. The imaging phantom assembly includes a phantom body and one or more coils. By arranging at least a part of the one or more coils in the phantom body, when the phantom body is subjected to geometric distortion correction imaging, the maximum distance between the coil and each part of the phantom body (for example, the coil and the central area of ​​the phantom body) is shorter, and the signal quality and accuracy of the phantom body acquired by the coil are improved, which can further improve the geometric distortion correction effect. The distance in one or more embodiments of this specification refers to the spacing along the main magnetic field direction of the device to be calibrated in the imaging space of the device to be calibrated (e.g., MRI device). Figure 2 In the illustrated embodiment, the main magnetic field direction of the MRI device can be represented by an arrow X, so when at least a portion of the one or more coils is disposed in the phantom body, the maximum distance between the coils and the various parts of the phantom body in the direction of the arrow X is shorter than when the coils are disposed at both ends of the phantom body in the X direction. In some embodiments, the requirements for different geometric distortion correction accuracies can be met by designing the number of coils, the relative positions of the coils and the phantom body, and the parameters of the coils.

[0019] In some embodiments, Figure 1 As shown, the imaging phantom assembly 100 includes a phantom body 110 and one or more coils 120, at least a portion of the one or more coils 120 is disposed in the phantom body 110. The phantom body 110 is an imaging phantom used to perform geometric distortion correction on a device to be corrected.

[0020] In some embodiments, the phantom unit 111 may have a regular shape (eg, spherical, cylindrical, cuboid, cube, etc.) or an irregular shape. Figure 1 and Figure 2 As shown, the phantom body 110 may include a plurality of phantom units 111, and the plurality of phantom units 111 may be arranged in a specific manner, and the shape of the phantom body 110 is related to the arrangement of the plurality of phantom units 111. Figure 2In the illustrated embodiment, a plurality of phantom units 111 may form a rectangular phantom body 110. For another example, a plurality of phantom units 111 may form a cylindrical phantom body 110. For another example, a plurality of phantom units 111 may form a spherical phantom body 110.

[0021] In some embodiments, the phantom unit 111 may be composed of any material capable of generating magnetic resonance under a magnetic field. In some embodiments, the phantom unit 111 may include a shell and a substance filled in the shell, and the substance filled in the shell is a material capable of generating magnetic resonance under a magnetic field, so that the atomic nuclei in the filling substance in the shell can enter a magnetic resonance state under the action of the magnetic field.

[0022] In some embodiments, the types of the phantom unit 111 may include a body phantom and a water phantom. The difference between a body phantom and a water phantom includes the different types of substances filled in the phantom. For example, when the filling substance is a liquid substance such as water, saline liquid, etc., the phantom unit 111 is a water phantom. For another example, when the filling substance is a material with a density similar to that of human tissue (e.g., plastic, epoxy resin, polyurethane, Mix D, cork, human bone, artificial bone, casting material, composite foam plastic, rubber, foam rubber, isocyanate rubber, thermoplastic material, Mix Dp, M3, E-type material and various other solids and liquids), the phantom unit 111 is a body phantom. For the convenience of description, some embodiments of this specification are described by taking water as the filling substance as an example. When performing geometric distortion correction imaging, after the RF signal is fed into the coil 120, the coil 120 can generate a magnetic field and generate RF energy, and the hydrogen atoms in the phantom unit 111 (i.e., the hydrogen atoms in the water) will receive the RF energy and enter an excited state (i.e., a magnetic resonance state). When the RF signal is stopped, the hydrogen atoms in the excited state will radiate energy outward, and the signal of the outward radiated energy is captured by the coil 120, thereby obtaining a magnetic resonance signal corresponding to each phantom unit 111. The magnetic resonance signal can be processed to obtain a correction image for geometric distortion correction imaging. The process of hydrogen atoms from entering the excited state to radiating energy outward can be called relaxation, so the signal of hydrogen atoms radiating energy outward obtained by the coil 120 can also be called a relaxation signal. For more details about geometric distortion correction imaging, please refer to Figure 6 The description of the embodiments thereof will not be repeated here.

[0023] In some cases, since at least a portion of one or more coils 120 is disposed inside the phantom body 110, it is equivalent to at least a portion of at least one coil 120 of the one or more coils 120 being located inside the imaging phantom assembly 100, thereby shortening the maximum distance between the phantom unit 111 located in the central area of ​​the phantom body 110 and the coil 120. The shortened distance means that when the phantom body 110 is subjected to geometric distortion correction imaging, the imaging effect of the corrected image is no longer limited by the way in which the coil 120 covers the phantom body 110, and the signal quality and accuracy of the phantom unit 111 at each position obtained by the coil 120 are higher, and the signal-to-noise ratio and resolution of the corrected image obtained by the correction imaging are higher, further improving the geometric distortion correction effect.

[0024] In some embodiments, the phantom body 110 may be a whole, and at least a portion of at least one coil 120 of the one or more coils 120 may be disposed in the phantom body 110. As an example only, the multiple phantom units 111 constituting the phantom body 110 are interconnected and inseparable, so that the phantom body 110 becomes an independent whole, and the one or more coils 120 may be embedded in the gap between two adjacent phantom units 111, or the one or more coils 120 may be wound around the one or more phantom units 111 to achieve the purpose of being disposed in the phantom body 110. In some embodiments, the one or more coils 120 may be completely disposed in the phantom body 110. For example, some of the coils 120 may be completely embedded in the phantom body 110, and other coils 120 may be disposed in the phantom body 110 along the main magnetic field direction (such as Figure 2 In some embodiments, a portion of one or more coils 120 may be disposed in the phantom body 110, and another portion may be disposed outside the phantom body 110. For example, a portion of a coil 120 may be embedded in the phantom body 110, and another portion may be disposed at at least one end of the phantom body 110 along the main magnetic field direction.

[0025] In some embodiments, in combination Figure 1 and Figure 2 As shown, the mold body 110 includes a plurality of mold body layers 112 , which are stacked, and the one or more coils 120 include one or more coil layers 121 , at least one of the one or more coil layers 121 is disposed between two adjacent mold body layers 112 .

[0026] The stacking of multiple model layers 112 means that multiple model layers 112 are stacked in the same position. Figure 2In the embodiment shown, it can be seen that the multiple model layers 112 are stacked in parallel layer by layer along their thickness direction. In some embodiments, the stacking direction of the multiple model layers 112 can be parallel to the main magnetic field direction, such as Figure 2 Indicated by the arrow X direction.

[0027] In this embodiment, since a plurality of phantom layers 112 are stacked, when at least one coil layer 121 is disposed between two adjacent phantom layers 112, it is equivalent to disposing at least one coil layer 121 inside the imaging phantom assembly 100 (the surfaces of the phantom layers 112 at both ends of the stacking direction can be regarded as the surfaces of the imaging phantom assembly 100), thereby shortening the maximum distance between each phantom layer 112 and the coil layer 121, and in particular shortening the distance between the phantom layer 112 and the coil layer 121 located in the central area of ​​the phantom body 110. In some cases, the shortening of the distance between the phantom layer 112 and the coil layer 121 located in the central area of ​​the phantom body 110 means that when the phantom body 110 is subjected to geometric distortion correction imaging, the imaging effect of the correction image is no longer limited by the covering method of the RF coil on the phantom, the signal quality and accuracy of the phantom layer 112 at each position acquired by the coil layer 121 are higher, and the signal-to-noise ratio and resolution of the correction image obtained by the correction imaging are higher, further improving the geometric distortion correction effect.

[0028] In some embodiments, the number of phantom layers 112 can be adjusted according to the size of the imaging space to be corrected. For example, when the size of the imaging space to be corrected is large, the number of phantom layers 112 can be increased or the thickness of a single phantom layer 112 can be increased to increase the size of the phantom body 110, thereby adapting to the larger imaging space. When the size of the imaging space to be corrected is small, the number of phantom layers 112 can be reduced or the thickness of a single phantom layer 112 can be reduced to reduce the size of the phantom body 110, thereby adapting to the smaller imaging space.

[0029] In some embodiments, when the model body 110 is formed by stacking a plurality of model layers 112, each model layer 112 may include a plurality of model units 111, and the plurality of model units 111 may be arranged in a specific manner into a layered structure to form the model layer 112. Figure 3 In the illustrated embodiment, a plurality of mold units 111 are distributed in a rectangular array to form a rectangular mold layer 112, and a plurality of rectangular mold layers 112 can be stacked into a cube or cuboid mold body 110. For another example, a plurality of mold units 111 are distributed in a circular matrix to form a circular mold layer 112, and a plurality of circular mold layers 112 can be stacked into a cylindrical mold body 110.

[0030] In some embodiments, Figure 2and Figure 3 As shown, the model layer 112 may include a base 1121, and the base 1121 is provided with a groove (not shown in the figure), and the multiple model units 111 of each model layer 112 may be arranged in the groove of the base 1121 of the model layer 112. In some embodiments, the multiple model units 111 in each model layer 112 are connected to each other. For example, in Figure 3 In the illustrated embodiment, the grooves in each row of the base 1121 are connected in sequence, so the mold units 111 disposed in the grooves are also connected to each other.

[0031] In some embodiments, each model layer 112 can be formed by splicing two or more structures. For example, the model layer 112 can include a first splicing layer and a second splicing layer, the first splicing layer includes a plurality of hemispherical model units 111, and the second splicing layer also includes a plurality of hemispherical model units 111. When the first splicing layer and the second splicing layer are spliced ​​along the stacking direction to form the model layer 112, the plurality of hemispherical model units 111 of the first splicing layer can be spliced ​​with the plurality of hemispherical model units 111 of the second splicing layer to form a plurality of model units 111 in the shape of a sphere.

[0032] In some embodiments, the number of coil layers 121 may be one, and in the stacking direction of the plurality of phantom layers 112, the distances between the coil layers 121 and the two ends of the imaging phantom assembly 100 are between 4:6 and 6:4, that is, the arrangement position of the coil layer 121 is close to the central area in the stacking direction of the phantom body 110, which can ensure that the distance between each phantom layer 112 and the coil layer 121 is shortened as much as possible when the number of coil layers 121 is the least, thereby improving the quality and accuracy of the relaxation signal collected by the coil layer 121. As an example only, the size of the phantom body 110 in the stacking direction is 500 mm (the size difference between the coil layer 121 and the phantom layer 112 in the stacking direction is large, so the influence of the coil layer 121 on the size of the phantom body 110 can be ignored), if the coil layer 121 is arranged at one end in the stacking direction of the phantom body 110, the farthest distance between the coil layer 121 and each phantom layer 112 is 500 mm (that is, the distance between the coil layer 121 and the other end in the stacking direction of the phantom body 110). When the setting position of the coil layer 121 is close to the central area of ​​the phantom body 110 in the stacking direction, for example, the distances between the two ends of the coil layer 121 and the phantom body 110 are 200 mm and 300 m respectively, the maximum distance between the coil layer 121 and each phantom layer 112 is 300 mm, so the distance between the coil layer 121 and each phantom layer 112 is not too far, which can ensure that the quality of the relaxation signal collected by the coil layer 121 meets the requirements of geometric distortion correction.

[0033] In some embodiments, when the number of coil layers 121 is one, the coil layer 121 can be arranged in the central area in the stacking direction of the phantom body 110. At this time, the distance between the coil layer 121 and the two ends of the phantom body 110 in the stacking direction is the same, and the farthest distance between the coil layer 121 and each phantom layer 112 is smaller (approximately equal to half the size of the phantom body 110 in the stacking direction), which can further ensure that the quality of the relaxation signal collected by the coil layer 121 meets the requirements of geometric distortion correction.

[0034] In some embodiments, the number of coil layers 121 may be multiple, and the multiple coil layers 121 may be spaced apart along the stacking direction of the multiple phantom layers 112. Spaced apart means that at least one phantom layer 112 is arranged between two adjacent coil layers 121. In some cases, by arranging multiple coil layers 121 at intervals between the two ends of the stacking direction of the phantom body 110, the maximum distance between the phantom layer 112 and the coil layer 121 can be further shortened, so that the relaxation signals of each phantom layer 112 can be more accurately collected. In addition, due to the increase in the number of coil layers 121, each coil layer 121 can obtain the relaxation signals of multiple phantom layers 112, that is, each phantom layer 112 is sampled multiple times, thereby improving the quality and accuracy of the magnetic resonance signal finally generated, and finally improving the quality of the corrected image.

[0035] In some embodiments, the plurality of coil layers 121 may be distributed at equal intervals along the stacking direction of the plurality of mold body layers 112. Equally spaced distribution means that the spacing between all two adjacent coil layers 121 is the same. For example, a mold body layer 112 is arranged between each two adjacent coil layers 121. In other embodiments, the plurality of coil layers 121 may be distributed at non-equal intervals along the stacking direction of the plurality of mold body layers 112, that is, the spacing between all two adjacent coil layers 121 is not exactly the same.

[0036] In some embodiments, the number of coil layers 121 is multiple, and in the stacking direction of the multiple phantom layers 112, the ratio of the distance between at least one coil layer 121 and the two ends of the imaging phantom assembly 100 is between 4:6 and 6:4, so as to shorten the distance between the phantom layer 112 and the coil layer 121 at different positions in the stacking direction, so that the coil layer 121 can more accurately collect the relaxation signal of each phantom layer 112. In some embodiments, when the number of coil layers 121 is multiple, one of the coil layers 121 can be arranged in the central area of ​​the phantom body 110 in the stacking direction, so as to further shorten the distance between the phantom layer 112 and the coil layer 121 at different positions in the stacking direction, so as to enable the coil layer 121 to more accurately collect the relaxation signal of each phantom layer 112.

[0037] In some embodiments, a receiving groove (not shown in the figure) is formed between two adjacent phantom layers 112, and a coil layer 121 is detachably inserted in the receiving groove. In some cases, by setting the coil layer 121 in a form that can be inserted in the receiving groove, the operator can freely combine the relative positions between the phantom layer 112 and the coil layer 121, and flexibly adjust the number and position of the coil layer 121. When correcting the geometric distortion of a large-sized imaging space, the number of phantom layers 112 of the imaging phantom assembly 100 is large. By freely combining the relative positions between the phantom layer 112 and the coil layer 121, and adjusting the number and position of the coil layer 121, the distance between the coil layer 121 and each phantom layer 112 can be flexibly adjusted to meet the requirements of different geometric distortion correction accuracy.

[0038] In some embodiments, the imaging phantom assembly 100 may further include an insert layer (not shown in the figure), the insert layer includes a phantom layer 112 combined with a coil layer 121, and the insert layer is detachably inserted in the accommodating groove. In some embodiments, the coil layer 121 can be combined with the phantom layer 112 by bonding, welding, snap-on connection, etc. In some cases, by combining the phantom layer 112 and the coil layer 121 into an insert layer, the coil layer 121 and the phantom layer 112 become a whole, and the number of coil layers 121 can be increased or decreased more conveniently, thereby reducing the complexity of the installation and removal of the coil layer 121. In addition, since the insert layer is formed by combining the phantom layer 112 and the coil layer 121, the phantom layer 112 of the insert layer can support and fix the coil layer 121, and there is no need to set an additional fixing structure between two adjacent phantom layers 112, which can effectively reduce the difficulty and cost of design and manufacturing and simplify the structure of the imaging phantom assembly 100.

[0039] In some embodiments, in combination Figure 2-Figure 3 As shown, the coil layer 121 includes a sub-coil 1211, and the projection of the area enclosed by the sub-coil 1211 in the stacking direction covers the projection of the model layer 112 in the stacking direction, so that the signals emitted by the multiple model units 111 of each model layer 112 during the relaxation process can be completely received by the coil layer 121, thereby obtaining high-quality and high-accuracy magnetic resonance signals.

[0040] In some embodiments, the coil layer 121 may include a plurality of sub-coils 1211, and the plurality of sub-coils 1211 are not connected to each other. Figure 3In the illustrated embodiment, the coil layer 121 includes four sub-coils 1211. In some embodiments, the projections of the regions enclosed by the multiple sub-coils 1211 in the stacking direction respectively correspond to a part of the projection of the phantom layer 112 in the stacking direction. In some cases, when the coil layer 121 includes multiple sub-coils 1211, since each sub-coil 1211 can simultaneously and independently acquire the relaxation signals of a part of the phantom units 111 of each phantom layer 112, and then generate corresponding magnetic resonance signals, the efficiency of the coil layer 121 in acquiring magnetic resonance signals is faster. In other cases, since the number of phantom units 111 of the phantom layer 112 covered by each sub-coil 1211 is smaller, the signal acquisition speed of a single sub-coil 1211 can be increased, thereby improving the signal acquisition efficiency of the coil layer 121.

[0041] In some embodiments, the coil layer 121 includes a support device (not shown in the figure), and the support device is used to wind the sub-coils 1211 to form the coil layer 121. In some embodiments, the support device can also provide a connection function to the phantom layers 112 on both sides of the coil layer 121. Exemplary support devices can include snap connection components, clamping components, support frames, etc.

[0042] In some embodiments, the projection of the region enclosed by the multiple sub-coils 1211 in the stacking direction covers the projection of the phantom layer 112 in the stacking direction, so that all the signals when the phantom units 111 of each phantom layer 112 relax can be completely received by the coil layer 121, thereby improving the accuracy and quality of the relaxation signals. Only as an example, in Figure 3 the illustrated embodiment, the projection of the phantom layer 112 in the stacking direction is rectangular, and the projections of the four sub-coils 1211 in the stacking direction are also rectangular and can cover the projection of the phantom layer 112 in the stacking direction after combination. However, not limited thereto, the projection of the phantom layer 112 in the stacking direction and / or the projections of the four sub-coils 1211 in the stacking direction can also be any other shapes, such as trapezoid, parallelogram, ellipse or circle, etc.

[0043] In some embodiments, there is no overlapping region between the enclosed region of one sub-coil 1211 among the multiple sub-coils 1211 and the enclosed region of at least another sub-coil 1211. For example, the coil layer 121 includes four sub-coils 1211, the shape of the enclosed region of each sub-coil 1211 is square, the four sub-coils 1211 are arranged in a "field" shape, and the projection in the stacking direction just covers the projection of the phantom layer 112 in the stacking direction.

[0044] In some embodiments, there is an overlapping region between the enclosed region of one sub-coil 1211 among the multiple sub-coils 1211 and the enclosed region of at least another sub-coil 1211. Only as an example, in Figure 3In the illustrated embodiment, the coil layer 121 includes four rectangular sub-coils 1211, and the enclosed area of ​​each sub-coil 1211 overlaps with the enclosed areas of the other three sub-coils 1211. However, this is not limited to this, and the coil layer 121 may also include sub-coils 1211 of other numbers and / or other shapes (e.g., trapezoidal, parallelogram, elliptical, or circular, etc.) as needed, and / or the enclosed area of ​​a part of the sub-coil 1211 overlaps with the enclosed area of ​​one or more other sub-coils 1211.

[0045] In some cases, when the enclosed area of ​​a sub-coil 1211 overlaps with the enclosed area of ​​at least another sub-coil 1211, it can be ensured that the corresponding phantom unit 111 near the edge of the enclosed area of ​​each sub-coil 1211 can be covered by the sub-coil 1211, so as to avoid missing the signal of the phantom unit 111 corresponding to the edge part, resulting in inaccurate generated magnetic resonance signals. Since the multiple sub-coils 1211 of each coil layer 121 are not connected, the overlapping part of the sub-coil 1211 and other sub-coils 1211 will not cause coupling problems, and the signal obtained by the sub-coil 1211 can be effectively prevented from being interfered. In addition, when the enclosed areas of multiple sub-coils 1211 overlap, it means that the multiple sub-coils 1211 will repeatedly collect the relaxation signals of the phantom units 111 in the overlapping areas, thereby improving the quality and accuracy of the relaxation signals of the phantom units 111 in this part, thereby improving the quality of the magnetic resonance signals.

[0046] In other embodiments, the shape of the enclosed area of ​​the sub-coil 1211 may not be limited to Figure 3 For example, the shape of the enclosed area may also include a regular or irregular shape such as a circle, a triangle, a polygon, a rectangle, etc. In one embodiment, the shapes of the enclosed areas of the plurality of sub-coils 1211 may be the same, such as Figure 3 As shown, the shapes of the enclosed areas of the four sub-coils 1211 are all rectangular. In some embodiments, the shapes of the enclosed areas of the multiple sub-coils 1211 may be different. In some embodiments, the shape of the enclosed area can be adjusted according to the projection shape of the model body layer 112 in the stacking direction to ensure that the projection of the enclosed area of ​​the coil layer 121 in the stacking direction can cover the projection of the model body layer 112 in the stacking direction. For example, when the projection shape of the model body layer 112 in the stacking direction is a square, in order to make the multiple sub-coils 1211 better cover the model body layer 112, the shape of the enclosed area of ​​the sub-coil 1211 can be a square or a rectangle. For another example, when the projection shape of the model body layer 112 in the stacking direction is a triangle, the shape of the enclosed area of ​​the sub-coil 1211 can be a triangle.

[0047] In some embodiments, Figure 4 and Figure 5 As shown, the phantom body 110 of the imaging phantom assembly 100 includes a plurality of first phantom layers 113 and one or more second phantom layers 114 . The plurality of first phantom layers 113 are stacked, at least one of the one or more second phantom layers 114 is disposed between two adjacent first phantom layers 113 , and the second phantom layer 114 is wound with a coil 120 .

[0048] The first mold body layer 113 in this embodiment is the same as or similar to the mold body layer 112 in the previous embodiment. The difference between the second mold body layer 114 in this embodiment and the mold body layer 112 in the previous embodiment is that the second mold body layer 114 is wound with a coil 120, that is, the second mold body layer 114 and the coil 120 are integrated into a whole by winding, and the second mold body layer 114 is detachably arranged between adjacent first mold body layers 113, instead of the coil 120 being detachably arranged between two adjacent mold body layers, and the number of coils 120 can also be easily increased or decreased, thereby reducing the complexity of the operation of installing and disassembling the coil layer 121.

[0049] like Figure 6 As shown, the present specification also provides an image generation method 600, which can be based on the imaging phantom assembly described in other embodiments of the present specification (for example, Figure 1-Figure 5 The imaging phantom assembly 100 of the magnetic resonance imaging device is used to perform geometric distortion correction on a specific space of the magnetic resonance imaging device. In some specific application scenarios, the image generation method 600 can perform geometric distortion correction on the imaging space of the MRI device. The image generation method 600 will be further introduced below by taking the application of the image generation method 600 to geometric distortion correction of the imaging space of the MRI device as an example. In some embodiments, the image generation method 600 can be manually performed by an operator or by a processor (not shown) of the MRI device, and the image generation method 600 includes:

[0050] Step 610: Acquire a magnetic resonance signal of the imaging phantom component under a radio frequency signal.

[0051] Wherein, obtaining the magnetic resonance signal of the imaging phantom component under the radio frequency signal refers to obtaining the magnetic resonance signal of the phantom body (for example, Figure 1-Figure 5In some embodiments, before acquiring the magnetic resonance signal of the phantom body under the radio frequency signal, the imaging phantom assembly can be placed in the imaging space of the device to be calibrated (for example, an MRI device), and the plane where the coil is located is perpendicular to the main magnetic field direction of the MRI device, so that the direction of the magnetic field generated after the coil receives the radio frequency signal is parallel to the main magnetic field direction, thereby improving the signal quality and accuracy of the coil acquisition. In some embodiments, the MRI device may include a moving component (not shown in the figure), the moving component may be connected to the imaging phantom assembly, and the processor may drive the moving component to control the imaging phantom assembly to move to the imaging space of the MRI device, and make the plane where the coil is located perpendicular to the main magnetic field direction of the MRI device.

[0052] In some embodiments, the size of the imaging phantom assembly can be adjusted according to the size of the imaging space to meet the correction requirements. In some embodiments, when the phantom body is a whole (for example, the phantom body 110 is directly composed of a plurality of phantom units 111 stacked), the size of the phantom body can be adjusted by increasing or decreasing the number of phantom units. In some embodiments, when the phantom body is composed of a plurality of phantom layers stacked, the size of the phantom body can be adjusted by increasing or decreasing the number of phantom layers. In some embodiments, the number of coils and the relative position between the coils and the phantom body can be adjusted according to the correction accuracy requirements. For example, when the correction accuracy requirements are high, the number of coils can be increased. When the correction accuracy requirements are low, the number of coils can be appropriately reduced. For another example, when the correction accuracy requirements are high, the coils can be set at a position close to the center area of ​​the phantom body.

[0053] In some embodiments, after the imaging phantom assembly is placed in the imaging space, the radio frequency signal transmitting assembly (not shown in the figure) of the MRI device is controlled to transmit a radio frequency signal to the coil of the imaging phantom assembly. After the radio frequency signal is fed into the coil, the coil can generate a magnetic field and generate radio frequency energy, and the hydrogen atoms in the phantom units of the phantom body receive the radio frequency energy and enter an excited state (i.e., a magnetic resonance state). At this time, the radio frequency signal transmitting assembly is controlled to stop working, and the hydrogen atoms in the multiple phantom bodies relax. The coil captures the signal of the relaxation of the hydrogen atoms and then generates a corresponding magnetic resonance signal.

[0054] In some embodiments, the number of coils is multiple, and the multiple coils are arranged at intervals relative to the phantom body along the main magnetic field direction (for example, the main magnetic field direction of the MRI device), and obtaining the magnetic resonance signal of the phantom body under the radio frequency signal may include: obtaining multiple sub-signals, each of the multiple sub-signals is obtained by a coil; and determining the magnetic resonance signal based on the multiple sub-signals. In this embodiment, when the imaging phantom assembly includes multiple coils, each coil can independently collect relaxation signals corresponding to the phantom unit at each position of the phantom body along the main magnetic field direction when the phantom unit radiates energy, that is, one coil can collect multiple relaxation signals corresponding to the phantom unit at each position of the phantom body along the main magnetic field direction when the phantom unit radiates energy. After the multiple relaxation signals collected by each coil are calculated and processed, a sub-signal can be obtained. In some embodiments, the multiple relaxation signals collected by a coil can be calculated and processed by summing, averaging, and variance, etc. to obtain the sub-signal corresponding to the coil. Since the positions of each coil are different, the radiation energy received by each coil from the phantom unit at each position of the phantom body along the main magnetic field direction is different, and thus the multiple relaxation signals collected by different coils are also different. For example, the farther the distance between the phantom unit located at a certain position of the phantom body and the coil along the main magnetic field direction, the weaker the phantom unit radiation energy received by the coil, and the weaker the relaxation signal intensity that the coil can collect. The closer the distance between the phantom unit located at a certain position of the phantom body and the coil along the main magnetic field direction, the stronger the phantom unit radiation energy received by the coil, and the stronger the relaxation signal intensity that the coil can collect. Therefore, when the setting position of the coil is different, the sub-signals acquired by the coil are also different, so when the imaging phantom assembly includes multiple coils, each coil can acquire a different sub-signal. Just as an example, the phantom body includes multiple phantom layers, one or more coils include one or more coil layers, at least one of the one or more coil layers is arranged between two adjacent phantom layers, and each coil layer can independently collect relaxation signals corresponding to the radiation energy of each phantom layer, that is, one coil layer can collect multiple relaxation signals corresponding to the radiation energy of each phantom layer. After calculation and processing, the multiple relaxation signals collected by each coil layer can obtain a sub-signal. In some embodiments, the multiple sub-signals acquired by the multiple coils can be processed by weighting and other methods to obtain a magnetic resonance signal. In some cases, when the imaging phantom assembly includes multiple coils, the multiple coils may repeatedly collect relaxation signals when one or more phantom units located at the same position of the phantom body along the main magnetic field direction (for example, multiple phantom units constituting the same phantom layer can be approximately regarded as being located at the same position in the main magnetic field direction) radiate energy, that is, multiple sampling is performed on one or more phantom units located at the same position of the phantom body along the main magnetic field direction, thereby improving the quality and accuracy of the magnetic resonance signal finally generated.

[0055] In some embodiments, when the imaging phantom assembly includes multiple coils, the RF signal transmitting assembly can be controlled to feed RF signals to the multiple coils at the same time. At this time, the multiple coils will simultaneously generate magnetic fields and transmit RF energy, and approximately synchronously obtain the sub-signals of the phantom units at various positions of the phantom body along the main magnetic field direction, thereby improving the image processing efficiency. In some embodiments, when the imaging phantom assembly includes multiple coils, the RF signal transmitting assembly can be controlled to feed RF signals to each coil in turn, and collect multiple sub-signals obtained by the multiple coils in turn. Since no other coil collects relaxation signals and obtains sub-signals at the same time during the process of collecting relaxation signals and obtaining sub-signals for each coil, interference with the collection of relaxation signals and the acquisition of sub-signals by the coil is avoided, and the problem of low quality of the generated magnetic resonance signal caused by the coupling of multiple coils can be effectively avoided. As an example only, the imaging phantom assembly includes a first coil, a second coil, and a third coil. The RF signal can be fed to the first coil first, and multiple first relaxation signals can be collected and the first sub-signals can be obtained through the first coil. Then the same RF signal is fed to the second coil, and multiple second relaxation signals can be collected and the second sub-signals can be obtained through the second coil. Finally, the same radio frequency signal is fed into the third coil, and the third relaxation signal is collected through the third coil to obtain a third sub-signal, and finally the magnetic resonance signal is determined based on the three sub-signals.

[0056] Step 620: Generate an image based on the magnetic resonance signal.

[0057] In some embodiments, after acquiring the magnetic resonance signal of the imaging phantom assembly, the magnetic resonance signal can be processed to generate an image. For example, the magnetic resonance signal can be subjected to k-space signal filling and signal post-processing to obtain an image (ie, a corrected image).

[0058] The beneficial effects of the imaging phantom assembly and the image generation method provided in the present specification include but are not limited to: (1) Since at least one of the one or more coils is arranged in the phantom body, the relaxation signals generated by the phantom layers at various positions in the phantom body can be well received by the coil, especially the distance between the coil and the one or more phantom units in the central area of ​​the phantom body is closer, the relaxation signals corresponding to the one or more phantom units received are more accurate, and the quality of the generated magnetic resonance signals is higher, thereby meeting the requirements of geometric distortion correction; (2) By arranging multiple coil layers at intervals between the two ends of the phantom body, it can be further ensured that the distance between the phantom layers at various positions and the coil layers is not too far, so that the signals corresponding to each phantom layer can be more accurately obtained. In addition, due to the increase in the number of coil layers and each coil being able to obtain signals from multiple phantom layers, there is cross data in the signals collected by the multiple coil layers, which is equivalent to improving the signal quality and accuracy of each phantom layer received, thereby further improving the quality and accuracy of the magnetic resonance signals, and ultimately improving the imaging quality; (3) By arranging the coil layers to be insertable in the container, the imaging quality can be improved. The form of the slot enables the operator to freely combine the relative positions between the phantom layer and the coil layer. When performing geometric distortion correction on a large-sized imaging space, the number of phantom layers in the imaging phantom assembly is large. By freely combining the relative positions between the phantom layer and the coil layer and adjusting the number of coil layers, the maximum distance between the coil layer and each phantom layer can be flexibly adjusted to meet the requirements of different geometric distortion correction accuracies; (4) By integrating the phantom layer and the coil layer into an insert layer, the number of coil layers can be more conveniently increased or decreased, thereby reducing the complexity of installing and removing the coil layer. In addition, since the insert layer is composed of the phantom layer and the coil layer, the phantom layer of the insert layer can support and fix the coil layer, and there is no need to set an additional fixing structure between two adjacent phantom layers, which can effectively reduce the difficulty and cost of design and manufacturing and simplify the structure of the imaging phantom assembly; (5) When the coil layer includes multiple sub-coils, since each sub-coil can simultaneously and independently receive the radiation energy of a part of the phantom of the phantom layer and then generate a corresponding signal, the coil layer can collect signals more efficiently.

[0059] The above description is only a preferred embodiment of this specification and is not intended to limit this specification. Any modifications, equivalent substitutions and improvements made within the spirit and principles of this specification should be included in the protection scope of this specification.

Claims

1. An imaging phantom assembly, characterized in that: include: The main body of the model; as well as One or more coils, at least a portion of at least one coil of the one or more coils is disposed within the phantom body.

2. The imaging phantom assembly according to claim 1, wherein: The mold body includes a plurality of mold body layers, and the plurality of mold body layers are stacked. The one or more coils include one or more coil layers, and at least one of the one or more coil layers is arranged between two adjacent mold body layers.

3. The imaging phantom assembly according to claim 2, characterized in that: There are multiple coil layers, and the multiple coil layers are spaced apart and distributed along the stacking direction of the multiple mold body layers.

4. The imaging phantom assembly according to claim 2, characterized in that: The number of the coil layer is one; in the stacking direction of the plurality of mold body layers, the ratio of the distance between the coil layer and the two ends of the mold body is between 4:6 and 6:

4.

5. The imaging phantom assembly according to claim 2, characterized in that: An accommodating groove is formed between two adjacent mold body layers, and the coil layer is detachably inserted into the accommodating groove.

6. The imaging phantom assembly according to claim 5, characterized in that: The imaging phantom assembly further includes an inserting sheet layer, wherein the inserting sheet layer includes the phantom layer combined with the coil layer, and the inserting sheet layer is detachably inserted into the containing groove.

7. The imaging phantom assembly according to claim 2, characterized in that: The coil layer includes a plurality of sub-coils.

8. The imaging phantom assembly according to claim 7, characterized in that: An enclosed area of ​​one of the sub-coils in the plurality of sub-coils of the coil layer overlaps with an enclosed area of ​​at least another sub-coil.

9. The imaging phantom assembly according to claim 1, wherein: The mold body includes a plurality of first mold body layers and one or more second mold body layers, the plurality of first mold body layers are stacked, and at least one of the one or more second mold body layers is arranged between two adjacent first mold body layers; wherein the coil is wound around the second mold body layer.

10. An image generation method, characterized in that: include: Acquire a magnetic resonance signal of the imaging phantom assembly according to any one of claims 1 to 8 under a radio frequency signal; An image is generated based on the magnetic resonance signals.

Images