Method for generating normalized correction coefficient of PET (positron emission tomography) imaging system

Abstract

The invention relates to a method for generating a normalized correction coefficient of a PET (positron emission tomography) imaging system, the same radioactive source is sequentially arranged at different axial positions of the PET system, the radioactive sources at different positions have overlapping parts in the axial direction, and the method comprises the following steps: splicing and fusing detection data at each position to obtain fusion data of a complete axial view; performing physical correction on the fused data to obtain conforming data; calculating an interlayer correction factor according to the number of intersection points of the response line and the radioactive source; performing axial interlayer correction on the coincidence data based on the interlayer correction factor to obtain correction data; and analyzing the correction data by adopting a model-based normalization correction method to obtain a normalization correction coefficient. The method has the advantages that the length of the radioactive source is not limited, the mechanical device of the short-axis PET system and the radioactive source are used for conducting normalization correction on the long-axis PET system, extra source use cost is not needed, and a complex mechanical device for fixing the radioactive source does not need to be redesigned.

Description

Technical Field

[0001] The invention relates to the technical field of medical imaging, and in particular to a method for generating a normalized correction coefficient of a PET imaging system. Background Art

[0002] Positron emission tomography (PET) is a high-end nuclear medicine imaging diagnostic device. In actual operation, radionuclides (such as 18 F. 11 C, etc.) to label the metabolites and inject the nuclides into the human body, and then use PET to perform functional metabolic imaging on the patient to reflect the status of life metabolic activities, thereby achieving the purpose of diagnosis.

[0003] The PET imaging system is usually a ring-shaped detection system containing tens of thousands of detection units. Affected by the geometric position and performance differences, such as crystal luminescence efficiency, crystal packaging, crystal coupling with PMT (photomultiplier tube) or silicon photomultiplier tube (SiPM), electronic system, different incident angles of photon pairs, etc., the detection efficiency of the detection unit is not consistent, resulting in its output not accurately reflecting the input photon beam intensity, which will inevitably introduce artifacts in the reconstruction process. In order to accurately model the detection system, the user must calibrate the detection efficiency of the detector in advance, that is, normalization correction.

[0004] Traditional normalization correction uses a uniform source placed at the center of the system (e.g., a rod source rotating at a constant speed, a cylindrical barrel source filled with FDG solution, a solid cylindrical Ge barrel source) to uniformly illuminate each detection unit, and uses a model-based normalization method to process the detection data. The intrinsic detection efficiency of the crystal, the normalization factor affected by system geometric factors, and the normalization factor related to the count rate are extracted from the data, and finally they are integrated into a normalization correction factor reflecting the overall detection efficiency, and the measurement data is normalized and corrected.

[0005] The current PET system has an increasingly larger axial field of view. In order to ensure that the radiation source can still evenly irradiate all detector crystals in the normalized test, the axial size of the radiation source is designed to be longer and longer, which greatly increases the weight and makes positioning and moving operations very difficult. In addition, the design and processing of the mechanical device that fixes the radiation source and the mechanical device that supports the stable and uniform reciprocating motion of the radiation source are extremely complex and expensive. Summary of the invention

[0006] Technical issues to be solved

[0007] In view of the above-mentioned shortcomings and deficiencies of the prior art, the present invention provides a method for generating a normalized correction coefficient of a PET imaging system, which solves the technical problems that the normalized correction currently requires additional source costs and the design and processing of the radiation source mechanical device is extremely complex and expensive.

[0008] Technical Solution

[0009] In order to achieve the above object, the main technical solutions adopted by the present invention include:

[0010] In a first aspect, the present invention provides a method for generating a normalized correction coefficient of a PET imaging system, wherein the same radiation source is placed successively at different axial positions of the PET system, and the radiation sources at different positions have overlapping portions in the axial direction, and the method comprises:

[0011] Modeling the PET acquisition process;

[0012] The acquisition time at each location is calculated based on the activity of the radioactive source at each location;

[0013] Collecting detection data based on the collection time at each location;

[0014] The detection data at each position are stitched and fused to obtain fused data covering the complete axial field of view;

[0015] Perform physical correction on the fused data to obtain conforming data;

[0016] The coordinates of the intersection of the response line and the radiation source are obtained according to the position information and system information recorded in each scan;

[0017] Determine whether the intersection exists according to the intersection coordinates, and calculate the interlayer correction factor according to the number of existing intersections;

[0018] Performing axial inter-layer correction on the conforming data based on the inter-layer correction factor to obtain corrected data;

[0019] The normalized correction coefficients were obtained by analyzing the correction data using a model-based normalized correction method.

[0020] Optionally, the PET acquisition process is modeled, including: the PET acquisition process is modeled as the following formula:

[0021]

[0022] in, represents the expected value of the detected data sinusoidal response line; r represents the variable index of the shortest distance between the response line and the radial center of the system, t represents the variable index of the azimuth of the response line, p represents the variable index of the axial fault plane to which the response line belongs, M represents the dimension of the sinusoidal r direction, S represents the dimension of the sinusoidal t direction, N represents the number of axial faults of the long-axis PET sinusoidal image, and the single quotation mark superscript represents the matrix transposition operation; x = [x 1 ,x 2 ,…,x j ,…,x Q ] ′ represents the unknown PET radioactivity distribution, j represents the number of the spatial point source, Q represents the size of the radioactivity distribution image space; A = [A jrtp ] is the system matrix, r=[r 111 ,r 112 ,…,r rtp ,…,r MSN ] ′ represents the average value of random noise and scattered noise; is the normalized correction factor matrix.

[0023] Optionally, the acquisition time at each location is calculated based on the activity of the radioactive source at each location, including:

[0024] According to the formula Calculate the acquisition time for each location, where T = [T 1 ,T 2 ,…,T i ,…T L ] represents the acquisition time of the radiation source at each position, α=[α 1 ,α 2 ,…,α i ,…α L ] represents the initial activity of the radioactive source at each location before collection, L represents the number of locations where the radioactive source is collected, and the subscript i represents the number of the location where the radioactive source is placed.

[0025] Optionally, the detection data at each position is spliced ​​and fused to obtain fused data covering the complete axial field of view, including:

[0026] According to the formula The detection data are spliced ​​and fused to obtain fused data covering the complete axial field of view, where: To fuse the data, For detection data.

[0027] Optionally, physical correction is performed on the fused data to obtain conforming data, including:

[0028] Perform attenuation correction, scatter correction, system response function correction, time drift correction, motion correction, noise correction and / or slice correction on the fused data to obtain conforming data.

[0029] Optionally, the coordinates of the intersection of the response line and the radiation source are obtained according to the position information and system information recorded in each scan, including:

[0030] The response line at any radial distance number r on the outermost plane of the PET system can determine a plane coordinate system O with the system axis. hoz , the h-axis coincides with the response line, the z-axis is always parallel to the system axis and falls on the inner surface of the detector, and the h-axis and the z-axis are perpendicular to each other;

[0031] In O hoz The position coordinates of the crystal pair corresponding to any response line on the graph are (z 1 ,h 1 ) and (z 2 ,h 2 ), the coordinate origin o is defined on the crystal surface, and the setting R scan Indicates the annular inner diameter of the PET system, R rho Indicates the actual distance between the response line numbered r and the center of the field of view, z 1 and z 2 Determined by the axial position of the crystal corresponding to the response line;

[0032] The formula for determining the response line according to the two crystal coordinates is:

[0033] h=k·z+b

[0034] in

[0035] The rotation trajectory of the radiation source at position i and the plane coordinate system O hoz There are two intersecting lines, which are two parallel straight lines in the coordinate system. The intersection line close to the positive direction of the coordinate axis h is defined as The intersection line close to the negative direction of the coordinate axis h is defined as

[0036] The intersection line at each position i is:

[0037]

[0038] in, Indicates intersection line and The axial coordinate near the negative end of the coordinate axis z, Indicates intersection line and The axial coordinate at the end close to the positive direction of the coordinate axis z;

[0039] Different radial positions, response lines and intersection lines and The axial intersection coordinates for:

[0040] Optionally, judging whether the intersection exists according to the intersection coordinates, and calculating the interlayer correction factor according to the number of existing intersections, includes:

[0041] if and All in If the response line is within the range, the number of intersections between the response line and the radiation source line is N. i rp =2; if and Only one in In the range, N i rp =1; if and None In the range, N i rp =0;

[0042] The number of intersections N between all response lines and the radiation source of the sine graph of the traversal detection data i rp , get N of the radiation source position i i rp ;

[0043] According to the formula Calculate the interlayer correction factor, where W rp is the inter-layer correction factor.

[0044] Optionally, performing axial inter-slice correction on the conforming data based on the inter-slice correction factor to obtain corrected data includes:

[0045] According to the formula Axial inter-slice correction is performed on the coincident data, where: To correct the data.

[0046] In a second aspect, the present invention provides a computer-readable storage medium having a computer program stored thereon, wherein when the program is executed, the method for generating a normalized correction coefficient of a PET imaging system as described in any one of the first aspects above is implemented.

[0047] In a third aspect, the present invention provides a storage device comprising a storage medium and a processor, wherein the storage medium stores a computer program, and when the program is executed by the processor, the method for generating a normalized correction coefficient of a PET imaging system as described in any one of the first aspects above is implemented.

[0048] Beneficial Effects

[0049] The beneficial effects of the present invention are as follows: a method for generating a normalized correction coefficient of a PET imaging system of the present invention, when collecting radioactive sources for generating normalized correction parameters, places radioactive sources at different positions of the long-axis PET system for multiple scanning, ensuring that the distance between the position of the radioactive source and the crystal surface is the same each time, and the radioactive sources placed multiple times can cover the complete axial field of view. The multiple collected radioactive source data are spliced ​​and fused, and the normalized correction parameters such as the detection efficiency, geometric correction factor and crystal interference correction factor of the long-axis PET system are analyzed using a model-based normalization method. This normalized correction method has no restrictions on the length of the radioactive source, and no longer requires a radioactive source that is sufficient to cover the entire axial field of view. The mechanical device and radioactive source of the traditional short-axis PET system can be used to perform normalized correction on the long-axis PET system, without requiring additional source costs or redesigning a complex mechanical device for fixing the radioactive source. BRIEF DESCRIPTION OF THE DRAWINGS

[0050] Figure 1 A schematic diagram of a flow chart of a method for generating a normalized correction coefficient of a PET imaging system provided by an embodiment of the present invention;

[0051] Figure 2 A schematic diagram of two position scans of a rotating rod source provided in an embodiment of the present invention;

[0052] Figure 3 The plane coordinate system O provided by the embodiment of the present invention hoz Schematic diagram;

[0053] Figure 4 A schematic diagram of the PET normalized correction coefficient crystal detection efficiency provided by an embodiment of the present invention;

[0054] Figure 5 A schematic diagram of a radial geometric correction factor of a PET normalization correction coefficient provided in an embodiment of the present invention;

[0055] Figure 6 A schematic diagram of a PET normalized correction coefficient crystal interference correction factor provided in an embodiment of the present invention. DETAILED DESCRIPTION

[0056] In order to better explain the present invention and facilitate understanding, the present invention is described in detail below in conjunction with the accompanying drawings through specific embodiments. Although exemplary embodiments of the present invention are shown in the accompanying drawings, it should be understood that the present invention can be implemented in various forms and should not be limited by the embodiments set forth herein. On the contrary, these embodiments are provided in order to enable a clearer and more thorough understanding of the present invention and to be able to fully convey the scope of the present invention to those skilled in the art.

[0057] First, refer to Figure 1 This embodiment provides a method for generating a normalized correction coefficient of a PET imaging system, wherein the same radiation source is placed successively at different axial positions of the PET system, and the radiation sources at different positions have overlapping parts in the axial direction. The method includes:

[0058] S1, Modeling the PET acquisition process.

[0059] S2, calculate the acquisition time at each location based on the activity of the radioactive source at each location.

[0060] S3, collecting detection data according to the collection time of each position.

[0061] S4, the detection data at each position are stitched and fused to obtain fused data covering the complete axial field of view.

[0062] S5, performing physical correction on the fused data to obtain conforming data.

[0063] S6, obtaining the coordinates of the intersection of the response line and the radiation source according to the position information and system information recorded in each scan.

[0064] S7, judging whether the intersection exists according to the intersection coordinates, and calculating the interlayer correction factor according to the number of existing intersections.

[0065] S8, performing axial inter-layer correction on the conforming data based on the inter-layer correction factor to obtain corrected data.

[0066] S9, using a model-based normalization correction method to analyze the correction data to obtain a normalized correction coefficient.

[0067] The present invention proposes a new method for generating normalized correction parameters, which collects radiation sources multiple times and splices data according to the position, and combines the model-based normalization method to obtain normalized correction parameters including detection efficiency, geometric correction factor, crystal interference correction factor, etc. Compared with the traditional method of using a long-axis radiation source that can cover the entire axial field of view to collect raw data for normalized correction, the method proposed in this article uses a short-axis radiation source to collect multiple times and then splice and fuse the data to generate a normalized correction coefficient. The advantages are very obvious: a universal short-axis radiation source can be used, and there is no need to design a long-axis radiation source for a long-axis PET system, which reduces the source cost. At the same time, there is no need to design and process mechanical devices that support the stability and uniform rotation of the long-axis source, which simplifies the equipment design and reduces the equipment cost.

[0068] Optionally, the PET acquisition process is modeled, including: the PET acquisition process is modeled as the following formula:

[0069]

[0070] in, Represents the expected value of the detected data sinogram response line LOR (line of response). The detected data sinogram response line LOR is scaled by three-dimensional coordinates: r represents the variable index of the shortest distance between the response line LOR and the radial center of the system, t represents the variable index of the azimuth of the response line LOR, p represents the variable index of the axial fault plane to which the response line LOR belongs, M represents the dimension of the sinogram in the r direction, S represents the dimension of the sinogram in the t direction, N represents the number of axial faults of the long-axis PET sinogram, and the single quote superscript represents the matrix transposition operation. x = [x 1 ,x 2 ,…,x j ,…,x Q ] ′ represents the unknown PET radioactivity distribution, j represents the number of the spatial point source, and Q represents the size of the radioactivity distribution image space. jrtp ] is the system matrix, which expresses the probability of a point source j at a spatial position in the PET system being detected by the response line LOR with coordinates (r, t, p) in mathematical form, including the geometric detection efficiency factor, photon detection point spread factor and photon attenuation factor in the photon pair detection process. The geometric detection efficiency is expressed as the probability that a photon pair reaches the surface of the detector pair, which depends on the change of the geometric detection efficiency of the detector pair relative to each pixel position; the photon point spread function is expressed as the response function of the detector crystal to the incident photon; and the photon attenuation is expressed as the probability that a photon pair decays before being detected. r = [r 111 ,r 112 ,…,r rtp ,…,r MSN ]′ Represents the average value of random noise and scattered noise. The normalization correction factor matrix is ​​used to correct the non-uniformity of the detector crystal detection efficiency, and the normalization parameters are independent of the flight time measurement results, so the flight time dimension is not considered. In order to reduce the requirements of the statistical data, the normalization correction factor evaluation usually adopts a model-based method. The purpose of the present invention is to optimize the process method for solving the normalization correction factor of the long-axis PET system.

[0071] Optionally, the acquisition time at each location is calculated based on the activity of the radioactive source at each location, including:

[0072] According to the formula Calculate the acquisition time for each location, where T = [T 1 ,T 2 ,…,T i ,…T L ] represents the acquisition time of the radiation source at each position, α=[α 1 ,α 2 ,…,α i ,…α L ] represents the initial activity of the radioactive source at each location before collection, L represents the number of locations where the radioactive source is collected, and the subscript i represents the number of the location where the radioactive source is placed.

[0073] The collected data is defined as L represents the number of locations where the radioactive source is collected, and the superscript i represents the number of the location where the radioactive source is placed. 1 ,T 2 ,…,T i ,…T L ] represents the acquisition time of the radioactive source at each location. Since the radioactive source decays irreversibly over time and the activity decreases continuously, the acquisition time at each location needs to be corrected in advance according to the activity. The starting activity of the radioactive source at each location is recorded before acquisition as α = [α 1 ,α 2 ,…,α i ,…α L ], calculate the acquisition time at each location, and ensure that the acquisition counts at each location are roughly the same.

[0074] Optionally, the detection data at each position is spliced ​​and fused to obtain fused data covering the complete axial field of view, including:

[0075] According to the formula The detection data are spliced ​​and fused to obtain fused data covering the complete axial field of view, where: To fuse the data, For detection data.

[0076] The detection data at each position are spliced ​​and fused to obtain a sinusoidal diagram of the detection data covering the complete axial field of view.

[0077] Optionally, physical correction is performed on the fused data to obtain conforming data, including:

[0078] Perform attenuation correction, scatter correction, system response function correction, time drift correction, motion correction, noise correction and / or slice correction on the fused data to obtain conforming data.

[0079] Considering that attenuation, dead time, scattering or random events are inevitable during the scanning process, it is necessary to perform physical correction on the combined data to try to eliminate the effects of attenuation, dead time, scattering or random events, so as to obtain more accurate combined data.

[0080] The axial inter-slice correction factor is calculated by estimating the correspondence of the axial crystal pairs when the radiation source is placed at different positions in the PET system. Taking the rotating rod source as an example, first, since the rod source is at a fixed radius R in the center of the field of view, rot Rotation, LOR has no difference at different circumferential angles, so the effect of angle on the axial interlayer correction factor is not considered. Secondly, for LOR at different radial positions, the distance between the radiation source and the crystal surface is different, so the axial compliance at different radial positions needs to be considered.

[0081] Optionally, the coordinates of the intersection of the response line and the radiation source are obtained according to the position information and system information recorded in each scan, including:

[0082] The response line at any radial distance number r on the outermost plane of the PET system can determine a plane coordinate system O with the system axis. hoz ,like Figure 3 As shown. The h-axis coincides with the response line, the z-axis is always parallel to the system axis and falls on the inner surface of the detector, the h-axis and the z-axis are perpendicular to each other, and the direction is customized;

[0083] In o hoz The position coordinates of the crystal pair corresponding to any response line on the graph are (z 1 ,h 1 ) and (z 2 ,h 2 ), the coordinate origin o is defined on the crystal surface, and h is set 1 =0; R scan Indicates the annular inner diameter of the PET system, R rho Indicates the actual distance between the response line numbered r and the center of the field of view, z 1 and z 2 Determined by the axial position of the crystal corresponding to the response line;

[0084] The formula for determining the response line according to the two crystal coordinates is:

[0085] h=k·z+b

[0086] in

[0087] The rotation trajectory of the radiation source at position i and the plane coordinate system O hoz There are two intersecting lines, which are two parallel straight lines in the coordinate system. The intersection line close to the positive direction of the coordinate axis h is defined as The intersection line close to the negative direction of the coordinate axis h is defined as

[0088] The intersection line at each position i is:

[0089]

[0090] in, Indicates intersection line and The axial coordinate near the negative end of the coordinate axis z, z i+ Indicates intersection line and The axial coordinate at the end close to the positive direction of the coordinate axis z;

[0091] For different radial positions, the above formulas can be combined to obtain the response line and the intersection line and The axial intersection coordinates for:

[0092] Optionally, judging whether the intersection exists according to the intersection coordinates, and calculating the interlayer correction factor according to the number of existing intersections, includes:

[0093] if and All in If the response line is within the range, the number of intersections between the response line and the radiation source line is N. i rp =2; if and Only one in In the range, N i rp =1; if and None In the range, N i rp =0;

[0094] The number of intersections N between all response lines and the radiation source of the sine graph of the traversal detection data irp , get N of the radiation source position i i rp ;

[0095] According to the formula Calculate the interlayer correction factor, where W rp is the inter-layer correction factor.

[0096] Set all position sources N i rp Sum them together to get the final axial interlayer correction factor.

[0097] Optionally, performing axial inter-slice correction on the conforming data based on the inter-slice correction factor to obtain corrected data includes:

[0098] According to the formula Axial inter-slice correction is performed on the coincident data, where: To correct the data.

[0099] Since the commonly used short-axis radiation source cannot fully fill the long-axis PET system, a segmented detection method is used to obtain normalized and corrected raw data of the complete axial field of view. This method will result in uneven collection of the number of axial inter-layer response lines LOR, and it is necessary to perform inter-layer count correction on the detection data according to various situations. The inter-layer correction factor is defined as W rp , the dimension of the correction factor is M×N.

[0100] Dividing the stitched data from the sinusoidal domain by the axial inter-slice correction factor can solve the data duplication problem caused by the axial intersection of the two acquisitions, thereby obtaining long-axis data with uniform counts.

[0101] For data Y sum-c′ Using the traditional model-based method, the normalized correction factor A is obtained by analysis. norm The corrected and spliced ​​detection data can be equivalent to the data of a rotating rod source that uniformly fills the long-axis axial field of view, and the traditional model-based normalization correction method can be directly applied.

[0102] In addition, the radiation source in the present application can be a rotating rod source, a hollow cylindrical source, and a plane source. When the radiation source is a rotating rod source, it is necessary to ensure that the rotation center and the rotation radius are the same.

[0103] Figure 2 Schematic diagram of a scan using a rotating rod source, acquired in two steps, with the two acquisitions having an overlapping area. Figure 4 To utilize the intrinsic detection efficiency obtained by the present invention, Figure 5 is the radial geometric correction factor obtained by using the present invention; Figure 6 is the crystal interference correction factor obtained by using the present invention.

[0104] Below is an annular inner diameter R scan =370mm, and the axial field of view length of 518mm is used as an example to perform a rod source normalization experiment on a PET system to illustrate the implementation steps of the present invention. The steps include:

[0105] Prepare a Ge rod source with a length of 300 mm and a diameter of 2 mm. The rod source is a common source for obtaining normalized correction parameters in the short-axis PET system. According to the length of the rod source, this experiment places it in two positions for collection.

[0106] Use a robotic arm to fix the rod source and rotate it at a constant speed around the central axis of the PET system's acquisition field of view. When placing the rotating rod source, it is necessary to ensure that the rod sources on both sides have the same rotation center, and the rotation radius is set to R. rot =330mm, and there is an overlap between the two acquisition positions in the axial field of view, ensuring that the two acquisitions cover the complete axial field of view of the PET system.

[0107] Place the rod source in the first position, that is, fix it to one side of the PET system by the cantilever, and ensure that the rod source covers both the axial edge of the PET system and the center of the axial field of view, and record The rod source rotates at a constant speed of 20 revolutions per minute within the PET system acquisition field of view, and the acquisition time T is set. 1 =129600s, the initial activity of the record collection is α 1 =300.73μmCi, collected data Y 1 .

[0108] After the first position is collected, remove the rod source and the fixing device, and place the rod source in the second position, that is, fix the rod source with the cantilever to the other side of the PET system, ensuring that the rod source covers the other axial edge of the PET system and the center of the axial field of view at the same time, and record It also moves uniformly inside the system at a rotation speed of 20 revolutions per minute. At this time, the initial activity collected is α 2 =299.56μmCi, according to the formula Calculate the total time T for the second acquisition 2 =130106s, collect data Y 2 .

[0109] After the two positions are collected, the collected data Y 1 and Y 2 Splice and fuse to get Y sum . Physical correction is performed to obtain the data Y sum-c .

[0110] Substitute the location information and system information recorded in the two scans into the formula:

[0111]

[0112] The coordinates of the intersection of the LOR and the radiation source at any radial position r can be obtained:

[0113]

[0114] R rho is the actual distance between the LOR numbered r and the center of the field of view. z can be determined based on the LOR position coordinates. 1 and z 1 . Traverse all LORs and judge and Is it in the range of [0,293.5]? Get N 1 rp ; judge again and Is it in the range of [230.0,518.0]? Get N 2 rp Based on the formula Calculate the accurate interlayer correction factor W rp .

[0115] Based on the formula For data Y sum -c Perform axial inter-slice correction. This way, you can get an equivalent full axial field of view consistent with data Y without the need for a long rod source that covers the entire axial range of the PET system. sum-c′ .

[0116] A model-based normalization correction method is used to sum-c′ By analyzing, the normalized correction coefficients such as crystal detection efficiency, radial geometry correction factor and crystal interference correction factor can be obtained. The results are shown in Figure 4 , Figure 5 and Figure 6 The new method for generating the normalized correction coefficient of the PET imaging system proposed by the present invention effectively simplifies the equipment design and reduces the source manufacturing cost.

[0117] In a second aspect, an embodiment of the present invention provides a computer-readable storage medium having a computer program stored thereon, wherein when the program is executed, the method for generating a normalized correction coefficient of a PET imaging system as described in any one of the first aspects above is implemented.

[0118] In a third aspect, an embodiment of the present invention provides a storage device, including a storage medium and a processor, wherein the storage medium stores a computer program, and when the program is executed by the processor, the method for generating a normalized correction coefficient of a PET imaging system as described in any one of the first aspects above is implemented.

[0119] It will be appreciated by those skilled in the art that embodiments of the present invention may be provided as methods, systems or computer program products. Therefore, the present invention may take the form of a complete hardware embodiment, a complete software embodiment, or an embodiment combining software and hardware. Furthermore, the present invention may take the form of a computer program product implemented on one or more computer-usable storage media (including but not limited to disk storage, CD-ROM, optical storage, etc.) containing computer-usable program code.

[0120] Obviously, those skilled in the art can make various modifications and variations to the present invention without departing from the spirit and scope of the present invention. Thus, if these modifications and variations of the present invention fall within the scope of the claims of the present invention and their equivalents, the present invention should also include these modifications and variations.

[0121] Although the embodiments of the present invention have been shown and described above, it is to be understood that the above embodiments are exemplary and are not to be construed as limitations of the present invention. A person skilled in the art may alter, modify, replace and modify the above embodiments within the scope of the present invention.

Claims

1. A method for generating a normalized correction coefficient for a PET imaging system, characterized in that: The same radioactive source is placed successively at different axial positions of the PET system, and the radioactive sources at different positions have overlapping parts in the axial direction. The method comprises: Modeling the PET acquisition process; The acquisition time at each location is calculated based on the activity of the radioactive source at each location; Collecting detection data based on the collection time at each location; The detection data at each position are stitched and fused to obtain fused data covering the complete axial field of view; Perform physical correction on the fused data to obtain conforming data; The coordinates of the intersection of the response line and the radiation source are obtained according to the position information and system information recorded in each scan; Determine whether the intersection exists according to the intersection coordinates, and calculate the interlayer correction factor according to the number of existing intersections; Performing axial inter-layer correction on the conforming data based on the inter-layer correction factor to obtain corrected data; The normalized correction coefficients were obtained by analyzing the correction data using a model-based normalized correction method.

2. The method for generating a normalized correction coefficient of a PET imaging system according to claim 1, characterized in that: Modeling the PET acquisition process includes: Modeling the PET acquisition process as follows: in, represents the expected value of the detected data sinusoidal response line; r represents the variable index of the shortest distance between the response line and the radial center of the system, t represents the variable index of the azimuth of the response line, p represents the variable index of the axial fault plane to which the response line belongs, M represents the dimension of the sinusoidal r direction, S represents the dimension of the sinusoidal t direction, N represents the number of axial faults of the long-axis PET sinusoidal image, and the single quote superscript represents the matrix transposition operation; x = [x1, x2, …, x j ,…,x Q ] ′ represents the unknown PET radioactivity distribution, j represents the number of the spatial point source, and Q represents the size of the radioactivity distribution image space; A = [A jrtp ] is the system matrix, r=[r 111 ,r 112 ,…,r rtp ,…,r MSN ] ′ represents the average value of random noise and scattered noise; is the normalized correction factor matrix.

3. The method for generating a normalized correction coefficient of a PET imaging system according to claim 2, characterized in that: The acquisition time for each location is calculated based on the activity of the radioactive source at each location, including: According to the formula Calculate the acquisition time for each location, where T = [T1, T2, …, T i ,…T L ] represents the acquisition time of the radiation source at each position, α=[α1,α2,Ω,α i ,…α L ] represents the initial activity of the radioactive source at each location before collection, L represents the number of locations where the radioactive source is collected, and the subscript i represents the number of the location where the radioactive source is placed.

4. The method for generating a normalized correction coefficient of a PET imaging system according to claim 3, characterized in that: The detection data at each position is stitched and fused to obtain fused data covering the complete axial field of view, including: According to the formula The detection data are spliced ​​and fused to obtain fused data covering the complete axial field of view, where: To fuse the data, For detection data.

5. The method for generating a normalized correction coefficient of a PET imaging system according to claim 4, characterized in that: Physical correction is performed on the fused data to obtain conforming data, including: Perform attenuation correction, scatter correction, system response function correction, time drift correction, motion correction, noise correction and / or slice correction on the fused data to obtain conforming data.

6. The method for generating a normalized correction coefficient of a PET imaging system according to claim 5, characterized in that: The coordinates of the intersection of the response line and the radiation source are obtained based on the position information and system information recorded in each scan, including: The response line at any radial distance number r on the outermost plane of the PET system can determine a plane coordinate system O with the system axis. hoz , the h-axis coincides with the response line, the z-axis is always parallel to the system axis and falls on the inner surface of the detector, and the h-axis and the z-axis are perpendicular to each other; In O hoz The crystal position coordinates corresponding to any response line on the crystal are (z1, h1) and (z2, h2). The origin o is defined on the crystal surface. R scan Indicates the annular inner diameter of the PET system, R rho It represents the actual distance between the response line with radial distance number r and the center of the field of view. z1 and z2 are determined according to the axial position of the crystal corresponding to the response line; The formula for determining the response line according to the two crystal coordinates is: h=k·z+b in The rotation trajectory of the radiation source at position i and the plane coordinate system O hoz There are two intersecting lines, which are two parallel straight lines in the coordinate system. The intersection line close to the positive direction of the coordinate axis h is defined as The intersection line close to the negative direction of the coordinate axis h is defined as The intersection line at each position i is: in, Indicates intersection line and The axial coordinate near the negative end of the coordinate axis z, Indicates intersection line and The axial coordinate at the end close to the positive direction of the coordinate axis z; Different radial positions, response lines and intersection lines and The axial intersection coordinates for:

7. The method for generating a normalized correction coefficient of a PET imaging system according to claim 6, characterized in that: Determine whether the intersection exists according to the intersection coordinates, and calculate the interlayer correction factor according to the number of existing intersections, including: if and Both If the response line is within the range, then the number of intersections between the response line and the radiation source line is N. i rp =2; if and Only one in In the range, N i rp =1; if and None In the range, N i rp =0; The number of intersections N between all response lines and the radiation source of the sine graph of the traversal detection data i rp , get N of the radiation source position i i rp ; According to the formula Calculate the interlayer correction factor, where W rp is the inter-layer correction factor.

8. The method for generating a normalized correction coefficient of a PET imaging system according to claim 7, characterized in that: Based on the inter-layer correction factor, the axial inter-layer correction is performed on the conforming data to obtain the corrected data, including: According to the formula r=1,…,M;t=1,…,S;p=1,…,N Perform axial inter-slice correction on the coincident data, where: To correct the data.

9. A computer-readable storage medium having a computer program stored thereon, characterized in that: When the computer program is executed by a processor, the method for generating a normalized correction coefficient of a PET imaging system as described in any one of claims 1 to 8 is implemented.

10. A storage device comprising a storage medium and a processor, wherein the storage medium stores a computer program, wherein: When the processor executes the computer program, the method for generating a normalized correction coefficient of a PET imaging system as described in any one of claims 1 to 8 is implemented.

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