PET detector imaging system and application thereof

Abstract

The invention relates to the technical field of positron emission tomography, in particular to a PET detector imaging system and application thereof, and the system comprises a shielding case, a scintillation crystal LYSO array, a silicon photomultiplier SiPM array, a PET circuit array, an FPGA main control board, a level conversion module, a Cameralink protocol conversion module and a PC upper computer. A model of a large-scale PET detector imaging system is established, a high-precision multi-channel PET circuit is adopted as a core component of the imaging system, a PET detector (a shielding cover, a scintillation crystal LYSO array and a silicon photomultiplier tube SiPM array) is modularized, an accurate annihilation position is coded by adopting a pulse time window conformity condition detection technology, and the annihilation accuracy is improved. The imaging system has the advantages of being extensible in scale, completely localized and high in sensitivity, can be effectively applied to the annular PET detector in the nuclear medicine field, and provides another implementation method and path for the large-scale medical diagnosis imaging technology.

Description

Technical Field

[0001] The present invention relates to the technical field of positron emission tomography (PET), and in particular to a PET detector imaging system and application thereof. Background Art

[0002] Positron emission tomography is a non-invasive imaging technique that can provide 3D tomographic images of the distribution of radioactive tracers in vivo. After compounds (choline, acetic acid, etc.) or analogs (fluorodeoxyglucose, i.e., FDG, etc.) containing radionuclides (such as 18F, 11C, 13N, 15O) required for human metabolism are injected into the human body, they will be distributed in the human body as they are metabolized. These radionuclides decay in the human body and produce positrons. After the positrons travel a short distance in the human tissue, they interact with electrons in the surrounding matter and annihilate, emitting two gamma photons (because the distance is short and not easy to measure, the location where the annihilation occurs can be approximately regarded as the location of the radionuclide). These highly penetrating gamma photons are detected, and the annihilation events are reconstructed using the detected signals, that is, the location distribution of the radionuclides is obtained. Tumor tissues generally show abnormal metabolic conditions. Therefore, the location of the tumor can be inferred by the concentration distribution of radionuclides in the human body in the form of compounds (or analogs) required for human metabolism. This is the principle of PET imaging for cancer diagnosis.

[0003] PET is one of the most advanced large-scale medical diagnostic imaging technologies today. PET equipment includes PET detectors and readout circuits. PET detectors include shielding covers, crystal arrays, and photoelectric conversion arrays or SiPMs. The principle of PET imaging is that a living organism located in a ring-shaped PET detector device and injected with an imaging agent will produce positrons due to the decay of the tracer. The positrons and negative electrons annihilate to emit two pairs of gamma rays moving in opposite directions at 180°. The crystals in the crystal array will convert the received gamma rays into photons, which are then converted into electrical signals through the photoelectric conversion array. The readout circuit reads the time and energy of the event based on the electrical signal, and then estimates the approximate location where the positron annihilates. By collecting a large amount of data, PET image reconstruction is achieved. The accuracy of the PET detector's readout of energy, time, and position information has a crucial impact on the quality of PET image reconstruction.

[0004] Traditional PET detector imaging systems have low levels of integration and localization, and their reliability, sensitivity, and positioning accuracy are also low. They cannot meet the needs of cutting-edge scientific research in nuclear medicine, and cannot solve the need for localization of PET detector imaging systems. Summary of the invention

[0005] In order to improve the problems existing in the prior art, that is, to improve the integration and localization degree of the PET detector imaging system, the present invention proposes a PET detector imaging system, including a shielding cover, a scintillation crystal LYSO array, a silicon photomultiplier tube SiPM array, a PET circuit array, an FPGA main control board, a level conversion module, a Cameralink protocol conversion module, and a PC host computer, wherein:

[0006] The shielding cover is used to be assembled outside the scintillation crystal LYSO array to reflect the visible light of the scintillation crystal LYSO array after being excited by high-energy gamma photons without loss;

[0007] LYSO array of scintillation crystals, used to release visible light after interacting with 511keV gamma photons;

[0008] Silicon photomultiplier tube SiPM array, used to convert received visible light photons into electrical signals, i.e. current pulse signals, through a photoelectric conversion array;

[0009] A PET circuit array is used to collect data on the current pulse signal of the silicon photomultiplier tube SiPM array and convert the electrical signal into energy and time signals;

[0010] FPGA main control board, used for image timing conversion of energy signals of PET detectors;

[0011] The level conversion module is used for level conversion to provide the required voltage for the entire system;

[0012] Cameralink protocol conversion module, used to perform data fitting and long-distance transmission of energy from 4 pairs of LVDS differential signals to the PET detector imaging system;

[0013] The PC host computer uses an image acquisition card to collect the position and energy information of the effect, and performs imaging and image reconstruction according to the frame rate set by the FPGA.

[0014] Compared with the prior art, the present invention has the following beneficial effects:

[0015] 1. The present invention proposes a PET detector (including a shielding cover + a scintillation crystal LYSO array + a silicon photomultiplier tube SiPM array) module, which is a completely domestically produced implementation scheme with high reliability and high sensitivity;

[0016] 2. The present invention can accurately collect and analyze the specific position where the crystal generates gamma rays by dividing the row and column structure of the SiPM array, which is of great significance for positioning and imaging analysis of the subsequent gamma ray generation position;

[0017] 3. The present invention uses the time window conditional detection technology to accurately encode the location where annihilation occurs, which can exclude the entry of many scattered photons and improve the detection sensitivity;

[0018] 4. The present invention uses an 18-bit high-precision ADC to digitize the energy collected by the PET circuit with high precision, and uses the FPGA main control board to perform Cameralink image conversion on the digitized energy signal, which has real-time analysis of the position and energy of the gamma rays generated by the crystal;

[0019] 5. The present invention uses four-channel serial LVDS low-voltage differential signal to transmit PET detector energy data, which can effectively offset external electromagnetic interference and is suitable for long-distance communication and data transmission in the PET detector imaging system. The system has good imaging quality and the transmission signal is not distorted;

[0020] 6. The present invention has the characteristics of low cost, scalability and complete localization;

[0021] 7. The present invention can be effectively applied to annular PET detector equipment in the field of nuclear medicine, providing another implementation method and path for large-scale medical diagnostic imaging technology, forming a complete positron emission tomography image, and providing strong support for medical diagnosis. BRIEF DESCRIPTION OF THE DRAWINGS

[0022] Figure 1 A system block diagram of a PET detector imaging system and an imaging method thereof provided in Embodiment 1 of the present invention;

[0023] Figure 2 A schematic diagram of the position of a PET detector imaging system provided in Example 1 of the present invention between a ring-shaped detector device, a bracket and a host computer;

[0024] Figure 3 A schematic structural diagram of a silicon photomultiplier tube SiPM array of a PET detector imaging system provided in Example 1 of the present invention;

[0025] Figure 4 A schematic diagram of pulse compliance conditions of a PET detector imaging system provided in Example 2 of the present invention;

[0026] Figure 5 A structural configuration relationship diagram of a PET detector module in a ring detector imaging device provided in Example 2 of the present invention;

[0027] Figure 6 A schematic diagram of the acquisition method of the x-axis and y-axis corresponding to the PET detector module "SiPM array + LYSO crystal array" in a ring detector imaging device provided in Example 2 of the present invention;

[0028] Figure 7 A schematic diagram of the timing of energy image conversion performed by the FPGA main control board of a PET detector imaging system provided in Example 1 of the present invention on a SiPM array;

[0029] Figure 8 A schematic diagram of a four-channel serial data output format of a Cameralink protocol conversion of a PET detector imaging system provided in Example 7 of the present invention;

[0030] Fig. 9 A schematic diagram of an energy transmission interface after Cameralink protocol conversion of a PET detector imaging system provided in Example 7 of the present invention;

[0031] Fig.10 A flow chart of signal acquisition and imaging conversion of a PET detector imaging system provided in Example 1 of the present invention;

[0032] Among them, 101, shielding cover; 102, scintillation crystal LYSO array; 103, silicon photomultiplier tube SiPM array; 104, PET circuit array; 105, FPGA main control board; 106, level conversion; 107, Cameralink protocol conversion; 108, PC host computer; 200, human body scanning channel; 201, collecting and receiving the person to be examined; 202, scanning channel wall; 203, PET detector module; 204, annular PET detector cover; 401, human body taking radioactive nuclides; 402, positrons; 403, a pair of PET detector modules; 501, field of view angle area; 502, human body channel. DETAILED DESCRIPTION

[0033] The following will be combined with the drawings in the embodiments of the present invention to clearly and completely describe the technical solutions in the embodiments of the present invention. Obviously, the described embodiments are only part of the embodiments of the present invention, not all of the embodiments. Based on the embodiments of the present invention, all other embodiments obtained by ordinary technicians in this field without creative work are within the scope of protection of the present invention.

[0034] The present invention proposes a PET detector imaging system, comprising a shielding cover, a scintillation crystal LYSO array, a silicon photomultiplier tube SiPM array, a PET circuit array, an FPGA main control board, a level conversion module, a Cameralink protocol conversion module, and a PC host computer, wherein:

[0035] The shielding cover is used to be assembled outside the scintillation crystal LYSO array to reflect the visible light of the scintillation crystal LYSO array after being excited by high-energy gamma photons without loss;

[0036] LYSO array of scintillation crystals, used to release visible light after interacting with 511keV gamma photons;

[0037] Silicon photomultiplier tube SiPM array, used to convert received visible light photons into electrical signals, i.e. current pulse signals, through a photoelectric conversion array;

[0038] A PET circuit array is used to collect data on the current pulse signal of the silicon photomultiplier tube SiPM array and convert the electrical signal into energy and time signals;

[0039] FPGA main control board, used for image timing conversion of energy signals of PET detectors;

[0040] The level conversion module is used for level conversion to provide the required voltage for the entire system;

[0041] Cameralink protocol conversion module, used to perform data fitting and long-distance transmission of energy from 4 pairs of LVDS differential signals to the PET detector imaging system;

[0042] The PC host computer uses an image acquisition card to collect the location and energy information of the effect, and transmits energy according to the frame rate set by the FPGA and reconstructs the image.

[0043] Example 1

[0044] like Figure 1 As shown, a system block diagram of a PET detector imaging system and an imaging method thereof includes a shielding cover 101, a scintillation crystal LYSO array 102, a silicon photomultiplier tube SiPM array 103, a PET circuit array 104, an FPGA main control board 105, a level converter 106, a Cameralink protocol converter 107, and a PC host computer 108. Among them:

[0045] ① Shielding cover 101: The scintillation crystal LYSO array 102 is provided with a shielding cover 101, which is a carbon fiber shielding cover, which shields the surrounding surfaces of the scintillation crystal LYSO array (except the bottom) to protect it from light, and can block the intake of external visible light. At the same time, the inner cortex is coated with reflective material, which can reflect the visible light of the scintillation crystal LYSO array 102 after being excited by high-energy γ photons to the greatest extent without loss. The bottom of the shielding cover 101 is not shielded, and receives the 511keV γ photons emitted by annihilation. In the present invention, there are many shielding covers 101, which are respectively assembled outside the scintillation crystal LYSO array, and only one is introduced here.

[0046] ② Scintillator crystal LYSO array 102: The full name of Scintillator crystal LYSO array 102 is "Yttrium Lutetium Silicate Scintillation Crystal", or Scintillator crystal LYSO for short. It has the characteristics of high light output, fast luminescence decay, large effective atomic number, high density, stable physical and chemical properties, no deliquesce, and high gamma ray detection efficiency. It is considered to be the best inorganic scintillator crystal material with comprehensive performance and is an ideal scintillator crystal for PET. The organism located in the annular PET detector and injected with imaging agent will produce positrons due to the decay of the tracer. The positrons and negative electrons annihilate and emit two gamma ray pairs moving in opposite directions at 180°. In the Scintillator crystal LYSO array 102, the received gamma rays are converted into photons. The absorption and blocking ability of LYSO for 511keV gamma photons is an important indicator of system sensitivity. The stronger the blocking ability, the higher the system sensitivity. Therefore, LYSO must have a large density and effective atomic number to shorten the attenuation distance of gamma photons as much as possible. Here, LYSO uses a density of 7.15g / cm 3 , relative light output 75 (% NaI (Tl)) γ-rays, decay time 40ns, crystal array 16*16, size 25*25*8mm 3 In the present invention, there are also a plurality of scintillation crystal LYSO arrays 102, and each scintillation crystal LYSO array 102 corresponds to a silicon photomultiplier tube SiPM array 103. In practical applications, the scale of the array can be expanded as needed, and only one is introduced here.

[0047] ③ Silicon photomultiplier tube SiPM array 103: The silicon photomultiplier tube SiPM array 103 converts the received photons into electrical signals and current pulse signals through the photoelectric conversion array. The SiPM detector in the present invention uses an 8×8 SiPM array for actual testing, specifically using the JARY-TP3050-8x8C model, breakdown voltage 25±0.2V, overvoltage 1~5V, number of channels: 8×8, photosensitive area: 3mm×3mm, array size: 25.38mm×25.38mm, pixel size: 50μm, number of pixels: 3364, with DF12B (3.0) -40DP-0.5V (86)) model connector. In the present invention, there are also many silicon photomultiplier tube SiPM arrays 103, and each silicon photomultiplier tube SiPM array 103 corresponds to a scintillation crystal LYSO array 102. In actual applications, the scale of the array can be expanded as needed. Here, only one is introduced.

[0048] ④PET circuit 104: A multi-channel high-precision PET circuit, referred to as PET circuit, is used here. The PET circuit 104 is packaged in BGA, and steel sheets and solder paste are used to perform ball and BGA welding at the bottom of the package, which can greatly save the space of the FPGA main control board, and more PET circuits 104 can be placed per unit area, which is convenient for integration and scale. The multi-channel high-precision PET circuit 104 can collect data on the current pulse signal of the silicon photomultiplier tube SiPM array 103, and convert the electrical signal into energy and time signals. The multi-channel high-precision PET circuit is a core component of the PET device and imaging system, and its performance is directly related to the performance of the system. In the present invention, there are also many PET circuits 104, and the PET circuit 104 corresponds to a set of PET detectors (shielding cover 101 + scintillation crystal LYSO array 102 + silicon photomultiplier tube SiPM array 103) below. In practical applications, the scale of the array can be expanded as needed, and only one is introduced here.

[0049] ⑤FPGA main control board 105: The core device of FPGA main control board 105 adopts Xilinx's spartan6, and the FPGA model is XC6SLX45T-3FGG484I for signal processing and control. The FPGA main control board converts the energy and time signals of 64 channels through the cameralink communication protocol. At the same time, the main control board also realizes the level conversion and provides power for each module. In the present invention, there is only one FPGA main control board 105, but many acquisition units (shielding cover 101 + scintillation crystal LYSO array 102 + silicon photomultiplier tube SiPM array 103 + PET circuit 104) are placed on the main control board, and they are all one-to-one corresponding. In practical applications, the scale of the array can be expanded as needed, and only one is introduced here.

[0050] ⑥Level conversion 106: Level conversion 106 performs level conversion to provide the required voltage for the entire system. The 5V voltage provided by the power supply is converted to 3.3V through the high-current LDO circuit LT1764EQ-3.3V, and then converted to 1.2V through the high-precision high-current LDO circuit LT3083. 3.3V and 1.2V are used to power the multi-channel high-precision PET imaging circuit and FPGA of the DUT board at the same time. 3.3V also needs to power the camera link conversion circuit.

[0051] ⑦ Cameralink conversion protocol 107: The present invention quantizes the energy and time signals transmitted back by each acquisition unit on the main control board into digital signals and corresponding positions according to the (M×N) array arrangement. Each unit will correspond to the size of the energy digital signal of the different row and column positions. Then, through the Cameralink communication protocol, the entire (M×N) array information is generated into an energy display image composed of frame rate, row signal, column signal and energy digital signal, and transmitted to the host computer through the video cable.

[0052] ⑧PC host computer 108: The energy image transmitted by the video cable is collected and reconstructed through the image acquisition card. The PC host computer 108 can determine the specific position and energy of the gamma rays received by the scintillation crystal LYSO array 102 in the PET detector through the transmitted information, and then estimate the approximate position where the positron annihilation occurs, and realize PET image reconstruction.

[0053] Example 2

[0054] like Figure 2 As shown, a schematic diagram of the position of a PET detector imaging system provided by an embodiment of the present invention between a ring-shaped detector device, a bracket and a host computer, that is, this embodiment proposes an application of a PET detector imaging system.

[0055] exist Figure 2 In the figure, the annular detector device and the bracket include: a human body scanning channel 200; a collection and receiving object to be examined 201; a scanning channel wall 202; a PET detector module 203 (the PET detector module includes a shielding cover, a scintillation crystal LYSO array, a silicon photomultiplier tube SiPM array, a PET circuit and an FPGA main control board); and an annular PET detector cover 204.

[0056] The scanning channel wall 202 arranged around the human body scanning channel 200 is used to fix the PET detector module 203. A plurality of PET detector modules 203 are placed between the scanning channel wall 202 and the annular PET detector cover 204 to collect and receive photon information emitted by the subject 201.

[0057] The host computer is placed next to the ring detector device, and the image data is transmitted through the video cable. The data converted by the Cameralink image protocol is transmitted to the host computer through the image acquisition card.

[0058] The present invention can be effectively applied to annular PET detectors in the field of nuclear medicine, providing another implementation method and path for large-scale medical diagnostic imaging technology, forming a complete positron emission tomography image, and providing strong support for medical diagnosis.

[0059] Example 3

[0060] This embodiment 1 provides a schematic structural diagram of a silicon photomultiplier tube SiPM array 103 of a PET detector imaging system, including:

[0061] The silicon photomultiplier tube SiPM array 103 converts the received photons into electrical signals and current pulse signals through the photoelectric conversion array; the top view of the connector structure of the silicon photomultiplier tube SiPM array 103 is shown in FIG. Figure 3 As shown above, two 40-pin connectors of model DF12B(3.0)-40DP-0.5V(86)) are used to connect to the FPGA main control board 105 at the top. The bottom is coupled to the scintillation crystal LYSO array 102 at the bottom. According to the needs, the contact surfaces of the two can be directly coupled or optical glue can be used for bonding and coupling.

[0062] The top view of the structure of the silicon photomultiplier tube SiPM array 103 is shown in FIG. Figure 3 As shown below, an 8×8 SiPM array is used, the first column of the x-axis is numbered from 1-1 to 8-1, and the first row of the y-axis is numbered from 1-1 to 1-8, forming an 8*8 energy image array, which corresponds to the energy display interface of the host computer.

[0063] The present invention only proposes a detailed description using an 8×8 SiPM array, but is not limited to this array size. It can be expanded to M×N (M represents the number of rows on the x-axis of the SiPM array, and N represents the number of columns on the y-axis of the SiPM array) according to the design requirements. The present invention can accurately collect the specific position of the gamma ray generated by the crystal based on the SiPM array, which has great reference significance for the subsequent positioning and energy analysis of the position of the gamma ray generated.

[0064] Example 4

[0065] like Figure 4 As shown, this embodiment provides a schematic diagram of pulse compliance conditions of a PET detector imaging system, including:

[0066] The annular detector imaging device is assembled using 32 PET detector modules 203, wherein the modules numbered 1-16 have a mirror image correspondence with the modules numbered 17-31 on a straight line, and are called a pair of PET detector modules 403;

[0067] The module x-axis and the module y-axis correspond to the x-axis and the y-axis of the top view of the SiPM array 103 structure, respectively, that is, in the annular detector imaging device, the x-axis is the horizontal axis of the section of the annular detector imaging device, the y-axis is the depth axis of the annular detector imaging device, and the z-axis is the vertical axis of the section of the annular detector imaging device. A human body 401 taking a radioactive nuclide decays in the body and releases a positron 402, which annihilates with an electron around it to produce a pair of gamma photons at 180° to each other;

[0068] The PET detector imaging system will use the conditional detection method to collect the data of the pair of gamma photons. In the PET circuit, a coincidence time window is set according to the collected time information, and the arriving pulse time signal is monitored to identify and judge whether it is a true coincidence event. The schematic diagram is shown in the figure. Figure 4 ;

[0069] The SiPM array number (1-1) in the PET detector module 1 and the SiPM array number (1-1) in the PET detector module 17 in the annular detector imaging device respectively detect two γ photons at 180° to each other generated by the annihilation of the positron 402. The PET detector module 1 and the PET detector module 17 detect a pulse signal generated by the γ photon. The pulse signal is timed through the set time window, and the pulse information is identified. Through the set time window, the energy of the signals entering the time window at the same time is compared. If the energy is the same, it is identified as a coincidence event and recorded. This is the coincidence identification process. The two pulses in the coincidence event correspond to a pair of γ photons at 180° to each other generated after the annihilation of the positron 402. The time window coincidence detection technology can exclude the entry of many scattered photons and improve the detection sensitivity.

[0070] Example 5

[0071] like Figure 5 As shown, this embodiment provides a structural configuration relationship diagram of a PET detector module in a ring detector imaging device.

[0072] The connection lines between the PET detector modules that meet the time window are obtained as response lines (LOR). Here, the PET detector module structure configuration diagram is to obtain the combination of all response line LORs. The detector module structure configuration is closely related to the structure of the annular detector device. There is a unique response line between two PET detector modules. However, in the process of image reconstruction, not every response line LOR collects data. The field of view (Field Of View, referred to as FOV) 501 of the annular detector imaging device is introduced here. In most annular detector imaging devices, the FOV diameter is set to about half of the detector ring diameter, because the response line LOR obtained by the PET detector module farther away from the center of the FOV is more seriously affected. In order to better obtain the response line LOR data, the diameter of the human body channel 502 adopts about half of the FOV diameter.

[0073] Encoding the addresses between PET detector modules is actually position encoding of the response lines LOR that meet the conditions in the time window. Each response line LOR passing through the field of view angle 501 area can reflect the position information of the annihilation event of the positron, determine the location where the annihilation occurs, that is, determine the position information of the image, so it is necessary to position encode all response lines LOR passing through the field of view angle 501. The encoding must satisfy that each response line LOR has one and only one corresponding encoding. Therefore, for all LYSO crystals on the annular detector ring, as long as the numbers of the two PET detector modules, the X-direction numbers and the Y-direction numbers of the two PET detector modules are determined, a response line LOR can be uniquely determined, and the location where the annihilation event occurs can also be uniquely determined.

[0074] It can be seen from the figure that PET detector module 1 corresponds to PET detector modules 13 to 22, and similarly, PET detector module 17 corresponds to PET detector modules 28 to 32 and 1 to 5. By analogy, all the corresponding relationships of the 32 PET detector modules can be found.

[0075] Example 6

[0076] like Figure 6 As shown, this is a schematic diagram of the x-axis and y-axis acquisition method corresponding to the PET detector module "SiPM array + LYSO crystal array" in a ring detector imaging device provided in this embodiment.

[0077] The 32 PET detector modules in the annular detector imaging device have a one-to-one correspondence, and the PET detector module "SiPM array + LYSO crystal array" in the present invention uses an 8×8 array. Therefore, the x-axis and y-axis corresponding to the 8×8 structure of the single PET detector module "SiPM array + LYSO crystal array" also have a corresponding correspondence. For example, in the figure, the PET detector module 1 and the PET detector module 17 are in a corresponding relationship, and the first column of the PET detector module 1 (x-axis) and each column of the PET detector module 17 (x-axis) are numbered from 1-1 to 8-1, that is, the first column number 1-1 of the PET detector module 1 (x-axis) and each column number of the PET detector module 17 (x-axis) from 1-1 to 8-1 are a corresponding relationship combination, and the same PET detector 1 module (x-axis) The remaining numbers of the first column are also the same and correspond to the numbers of each column of the PET detector module 17 (x-axis) from 1-1 to 8-1. By analogy, the 2nd to 8th columns of PET detector module 1 (x-axis) and each column number of PET detector module 17 (x-axis) are also completely corresponding. Here, two corresponding PET detector module x-axis columns are used as examples, and the same can be done with y-axis rows, which will not be explained one by one here.

[0078] It can be seen from the figure that the number (1-1) of the PET detector module 1 and the PET detector modules 17 from (1-1) to (8-1) are possibly corresponding, so as long as the number is recorded until a response line LOR where a positron annihilation event occurs is found, and the numbers of the two PET detector modules are determined in Example 5, and the column numbers in the X direction and the row numbers in the Y direction of the two PET detector modules are recorded together, a complete position code, that is, position information, is obtained.

[0079] Example 7

[0080] like Figure 7 This embodiment provides a schematic diagram of the timing of energy image conversion performed by the FPGA main control board 105 of a PET detector imaging system on the SiPM array 103.

[0081] The FPGA main control board 105 determines whether the information collected by the PET circuit 104 meets the conditions through the time window, determines which channel collects which position and has the effect, and matches the position coding information with the collected energy information. Here, the FPGA main control board 105 uses the 18-bit 8-channel AD7608BSTZ to digitize the energy analog signal collected by the PET circuit 104. The AD7608BSTZ uses 8-way synchronous sampling input, true bipolar analog input, and 5V single analog power supply. According to the structural schematic diagram of the SiPM array 103, the FPGA main control board 105 performs image conversion on it, and generates the accompanying clock CLK, frame signal FRAME, row signal ROW, row valid DAT, image data SD0 (LSB) to image data SD17 (MSB).

[0082] AD7608BSTZ digitizes the energy signals of the 8 channels collected, and the FPGA main control board 105 performs image conversion on the energy signals of the 8 channels respectively. The SiPM array 103 adopts an 8*8 array. For example, if the first row is encoded as (1-1), (1-2), (1-3), ... (1-8) row pixels, the 8 row pixels of the first row are digitized using 18 bits, and the 18 bits are encoded as bit0, bit1, bit2, ... bit17, corresponding to the image data SD0, SD1, SD2, ... SD17. Here, the image data is serially input, and the image data SD0 input sequence is the bit0 data of the first row (1-1), (1-2), (1-3), ... (1-8), followed by the bit0 data of the second row (1-1), (1-2), (1-3), ... (1-8), and so on, until the bit0 data of the eighth row (1-1), (1-2), (1-3), ... (1-8) are all transmitted. Similarly, the image data SD1, SD2, ... SD17 also transmits bit1, bit2, ... bit17 in serial. Here, bit0, bit1, bit2, ... bit17 of the same row of pixels are transmitted in parallel.

[0083] When the first row of pixel image data is transmitted, the FPGA will generate a high-level row signal ROW and row valid DAT. After the first row of data is completely transmitted, a low-level row signal ROW and row valid DAT are generated. And so on, until the eighth row of pixel image data is transmitted and the row signal ROW and row valid DAT are pulled low. When the first row of pixel image data of the first row is transmitted, the FPGA will generate a high-level frame signal FRAME. After all the pixel image data of the 8×8 array is transmitted, the frame signal FRAME is pulled low. The FPGA generates a companion clock CLK for each pixel data transmission until the image is completely transmitted.

[0084] When transmitting data, the present invention uses the rising edge of the accompanying clock CLK signal to collect data. By performing timing operations on FRAME, ROW, CLK and DAT, the design selects an image resolution of M×N (M represents the number of rows of the SiPM array, and N represents the number of columns of the SiPM array, which is selected according to design requirements. Here, an 8×8 SiPM array is selected for illustration) and a data output of a frame rate F (unit Hz). In order to achieve the F frame rate, it is necessary to determine the number of row valid pixel clock cycles A, the number of row invalid clock cycles B, and the number of frame invalid clock cycles C. The frequency of the accompanying clock CLK is F clk , the specific implementation is as follows:

[0085]

[0086] Finally, the FPGA main control board 105 sends the 18-bit parallel data signal and FRAME, ROW, CLK and DAT generated according to the F frame rate to the Cameralink interface.

[0087] The present invention numbers the SiPM array to form an M×N (here 8×8) image array, uses the 18-bit high-precision AD7608BSTZ to digitize the energy collected by the PET circuit, and uses the FPGA main control board to perform Cameralink image conversion on the digitized energy signal. Thus, the energy information of the specific position of the gamma ray generated by the crystal is accurately transmitted, and the position and energy analysis of the gamma ray generated by the crystal is real-time.

[0088] Example 8

[0089] As shown in Table 1, the corresponding relationship of the output data formats of the Cameralink protocol converter 107 of the PET detector imaging system in this embodiment.

[0090] Table 1 Output data format of Cameralink protocol conversion

[0091]

[0092] Cameralink is used in medical imaging, and can transmit high-definition medical images in real time, providing doctors with accurate diagnostic evidence. Cameralink is a video signal output serial communication protocol that uses low-voltage differential signal LVDS transmission. In the standard CameraLink interface, the image acquisition card port definition is shown in the leftmost column of the table, and the digital image camera signal is divided into: video image data (PortA-PortH) and enable signals (Fval, Lval, Dval and Spare). The CameraLink interface standard is a specification for digital cameras and image acquisition cards, allowing the two to transmit multi-bit image data and enable signals (row signal, frame signal, row valid signal and spare signal) at a high-speed rate. The definition of the interface standard 28-bit input of the image acquisition card is shown in the middle column of the table. Cameralink has three connection modes: BASE mode (port Port A~Port C), Medium mode (port Port A~Port F) and Full mode (port Port A~Port H). The present invention selects the Cameralink interface configuration as Base. In Base mode, 18-bit and 24-bit RGB configurations can be used. For the entire system, 18-bit with a total bit width of 28 bits under the Base configuration will be used to meet the input requirements of 18-bit pixel data.

[0093] The present invention digitizes the energy signals of the 8-channel PET circuit collected by the 18-bit AD7608BSTZ. The 18-bit numbers are respectively encoded as bit0, bit1, bit2, ... bit17, which correspond one to one with the definition of the interface standard 28-bit input of the image acquisition card, and the bits without corresponding bits are input as 0 for completion. The corresponding relationships are (bit0~IN0), (bit1~IN1), (bit2~IN2), (bit3~IN3), (bit4~IN4), (bit5~IN6), (bit6~IN27), (bit7~IN5), (bit8~IN7), (bit9~IN8), (bit10~IN9), (bit11~IN12), (bit12~IN13), (bit13~IN14), (bit14~IN10), (bit15~IN11), (bit16~IN15), (bit17~IN18), (0~IN19), (0~IN20), (0~IN21), (0~IN22), (0~IN16), (0~IN17). The FPGA main control board 105 transmits the generated frame signal FRAME, row signal ROW and row valid signal DAT to the image acquisition card FVAL, LVAL and DVAL ports, and at the same time, the FPGA main control board 105 inputs the spare signal as 0 and transmits it to the image acquisition card Spare port. These four signals correspond to the definition of the interface standard 28-bit input of the image acquisition card one by one, and the corresponding relationship is (ROW~IN24), (FRAME~IN25), (DAT~IN26), (0~IN23). Finally, according to the requirements of TxCLK Out / TxCLK In, the FPGA main control board 105 generates the corresponding clock.

[0094] Example 9

[0095] like Figure 8 , this embodiment provides a schematic diagram of the four-channel serial data output format of the Cameralink protocol converter 107 of a PET detector imaging system.

[0096] The Cameralink protocol is a serial communication protocol specifically for the field of machine vision applications. It uses low voltage differential signaling (LVDS) for data transmission and communication. Here, the PET detector imaging system is converted to the Cameralink protocol according to the Cameralink protocol. The present invention uses low voltage differential signaling (LVDS) for data transmission and communication. Four pairs of LVDS differential signals are used to perform data fitting and transmission on the digital signal after quantization of the PET detector energy collection. The four pairs of LVDS differential signals are TxOUT3 / RxIN3, TxOUT2 / RxIN2, TxOUT1 / RxIN1, and TxOUT0 / RxIN0. The coupling relationship between the frame signal FRAME, the row signal ROW, the row valid DAT, and the pixel data (bit0 to bit17) and the four pairs of LVDS differential signals is shown in Example 4. In the present invention, the FPGA main control board 105 also generates a TxCLK OUT / RxCLK IN clock signal. Note that the clock signal here is not the same signal as the accompanying clock CLK on the left. The TxCLK OUT / RxCLK IN clock signal here is an LVDS differential parallel clock with a duty cycle of 4:3, that is, the high level occupies 4 and the low level occupies 3. It is used for data alignment at the receiving end. After 7 times the frequency, it can be used for the image acquisition card to collect differential serial data. The high and low levels are used to transmit the data of four pairs of differential pairs TxOUT3 / RxIN3, TxOUT2 / RxIN2, TxOUT2 / RxIN2, and TxOUT1 / RxIN1 in parallel. After the TxCLK OUT / RxCLK IN clock signal starts to rise, the four pairs of differential pairs TxOUT(3~0) / RxIN(3~0) (i.e., the digital signals of the corresponding PET detector energy) TxIN5(bit7), TxIN20(0), TxIN9(bit10), and TxIN1(bit1) are transmitted respectively. The rising edge, falling edge, high and low levels of the clock signal and the parallel data must be strictly aligned. And so on, no more elaboration is needed here, as shown in the example diagram.

[0097] The present invention uses four-channel serial LVDS low-voltage differential signals to transmit PET detector energy data, because data transmission is usually affected by external interference, which affects two signal lines at the same time. The differential signal transmits data through the difference between the two to effectively offset the external electromagnetic interference. At the same time, common-mode noise is also interference that appears simultaneously on the signal line, and the differential signal only focuses on the potential difference between the two lines. Therefore, the differential signal can also effectively suppress common-mode noise, which is suitable for long-distance communication and data transmission in the PET detector imaging system. The energy transmission signal is not distorted, and the system imaging quality is good.

[0098] Example 10

[0099] like Fig. 9 , this embodiment provides a schematic diagram of the energy transmission interface after the Cameralink protocol conversion of a PET detector imaging system.

[0100] The PET detector is converted through the Cameralink protocol and then transmitted to the host computer for energy information transmission. The energy transmission interface can display the energy of the silicon photomultiplier tube SiPM array in Example 3 respectively. According to the array scale of the PET detector, the "shielding cover + scintillation crystal LYSO array + silicon photomultiplier tube SiPM array" PET detector generates energy display interfaces such as PET detector 1, PET detector 2, and PET detector 3 respectively. For example, in PET detector 1, each 8×8 SiPM array is displayed as a pixel in the energy transmission screen. According to the required energy transmission interface array such as M×N, n 8×8 SiPM arrays can be selected for display, and their calculation formula is M×

[0101] N=n×(8×8).

[0102] Embodiment 11

[0103] like Fig.10 , this embodiment provides a flow chart of signal acquisition conversion imaging of a PET detector imaging system. It is used to realize the acquisition of energy and time signals of the PET detector, use the pulse signal that meets the conditions of the acquisition time window to determine the location information of the positron annihilation event, and use the energy information to perform Cameralink image protocol conversion. This embodiment includes the following steps:

[0104] Step 1101, the scintillation crystal LYSO array interacts with the gamma photons and releases the energy in the form of visible light photons. As in Example 1, the scintillation crystal LYSO array 102 interacts with the 511keV gamma photons in the shielding cover 101 and releases the energy in the form of visible light.

[0105] Step 1102, the silicon photomultiplier tube SiPM array converts the visible light signal collected in the corresponding crystal into an electrical signal. As in Example 1, the silicon photomultiplier tube SiPM array 103 converts the received visible light photons into electrical signals and current pulse signals through the photoelectric conversion array.

[0106] Step 1103, the multi-channel PET circuit array converts the electrical signal corresponding to the SiPM array into energy and time signals. As in Example 1, the multi-channel PET circuit 104 can collect data on the current pulse signal of the silicon photomultiplier tube SiPM array 103 and convert the electrical signal into energy and time signals.

[0107] In step 1104, the FPGA performs a time window conformity judgment on the signal input by the PET circuit array, and records the position information where the effect occurs, digitizes the energy signal corresponding to the position, and performs Cameralink protocol conversion. As in Example 1, the FPGA main control board 105 judges which channel collects which position and has an effect based on the information collected by the multi-channel PET circuit 104, and then locates it according to the structural diagram of the SiPM array 103 and performs position encoding. The specific process is shown in Examples 2 to 6.

[0108] The FPGA main control board 105 is powered by the level conversion module 106, and an 18-bit 8-channel ADC is selected to digitize the energy signal at the position, and image timing conversion is performed on it to generate the accompanying clock CLK, frame signal FRAME, row signal ROW, row valid DAT, and image data. The specific process of image timing conversion is shown in Example 7.

[0109] The digitized image data (energy) and the generated four control signals are matched one by one according to the requirements of the Cameralink protocol conversion 107. The specific process of the matching relationship and the output data format is shown in Example 8.

[0110] The corresponding Cameralink data format is transmitted and communicated using the anti-interference and low-noise LVDS low-voltage differential signal. Cameralink protocol conversion 107 uses four pairs of LVDS differential signals to perform data fitting and long-distance transmission of the energy of the PET detector imaging system. The specific process of the four-channel serial data output format is shown in Example 9.

[0111] Step 1105, the host computer image acquisition card acquires the energy signal transmitted at the frame rate set by the FPGA, obtains the position information of the corresponding energy, and finally reconstructs the image. The specific process of the FPGA main control board 105 setting the frame rate is shown in the calculation formula in Example 7. The PC host computer 108 uses the image acquisition card to transmit an energy signal transmitted by the long line at a certain frame rate, and then determines the specific position and energy size of the scintillation crystal LYSO array 102 in the PET detector to receive the gamma ray, and then estimates the approximate position where the positron annihilates, and reconstructs the data by collecting a large amount of PET position and energy image data, and draws a complete positron emission tomography image, which provides strong support for medical diagnosis.

[0112] The present invention builds a model of a large-scale PET detector imaging system, adopts a high-precision multi-channel PET circuit as the core component of the imaging system, and adopts the Cameralink protocol for energy signal acquisition and conversion and energy interface transmission, providing another implementation method and path for large-scale medical diagnostic imaging technology. The present invention has the characteristics of scalable PET detector, high imaging accuracy, strong anti-interference ability of image signals, and acquisition of a large amount of detector data for image reconstruction, and can be effectively applied to medical annular PET detectors.

[0113] The above embodiments further illustrate the purpose, technical solutions and advantages of the present invention in detail. It should be understood that the above embodiments are only preferred implementation modes of the present invention and are not intended to limit the present invention. Any modifications, equivalent substitutions, improvements, etc. made to the present invention within the spirit and principles of the present invention should be included in the protection scope of the present invention.

Claims

1. A PET detector imaging system, characterized in that: It includes a shielding cover, a scintillation crystal LYSO array, a silicon photomultiplier tube SiPM array, a PET circuit array, an FPGA main control board, a level conversion module, a Cameralink protocol conversion module, and a PC host computer, including: The shielding cover is used to be assembled outside the scintillation crystal LYSO array to reflect the visible light of the scintillation crystal LYSO array after being excited by high-energy gamma photons without loss; LYSO array of scintillation crystals, used to release visible light after interacting with 511keV gamma photons; Silicon photomultiplier tube SiPM array, used to convert received visible light photons into electrical signals, i.e. current pulse signals, through a photoelectric conversion array; A PET circuit array is used to collect data on the current pulse signal of the silicon photomultiplier tube SiPM array and convert the electrical signal into energy and time signals; FPGA main control board, used for image timing conversion of energy signals of PET detectors; The level conversion module is used for level conversion to provide the required voltage for the entire system; Cameralink protocol conversion module, used to perform data fitting and long-distance transmission of energy from 4 pairs of LVDS differential signals to the PET detector imaging system; The PC host computer uses an image acquisition card to collect the location and energy information of the effect, and transmits energy according to the frame rate set by the FPGA and reconstructs the image.

2. A PET detector imaging system according to claim 1, characterized in that: The shielding cover is a carbon fiber shielding cover, which has a through-hole structure and the inner skin is coated with reflective material.

3. A PET detector imaging system according to claim 1, characterized in that: A shielding cover is mounted on the outer surface of a scintillation crystal LYSO array. One end of the scintillation crystal LYSO array receives γ photons and interacts with them, and the other end is coupled with the scintillation crystal LYSO array. A shielding cover, a scintillation crystal LYSO array and a silicon photomultiplier tube SiPM array form a PET detector. The PET circuit array is arranged on an FPGA main control board. A PET detector is connected to the FPGA main control board via a connector.

4. An application of a PET detector imaging system, characterized in that: A detection channel is formed by using a scanning channel wall and a PET detector cover, and a plurality of PET detector imaging systems according to claim 1 are installed in the annular area between the scanning channel wall and the PET detector cover, and a PET detector imaging system is used to collect photon information emitted by a person to be detected in the detection channel.

5. The use of a PET detector imaging system according to claim 4, characterized in that: When two PET detector imaging systems located on the diameter of the circle passing through the annular area are regarded as a pair, the pair of imaging systems respectively detect two γ photons that are 180° to each other, and the pulse signal is timed and the pulse information is identified through the set time window. Through the set time window, the energy of the signals that enter the time window at the same time is compared. If the energy is the same, it is a valid event. Only the visible light photons of the valid event are converted into electrical signals through the photoelectric conversion array to perform subsequent operations.

6. The use of a PET detector imaging system according to claim 4, characterized in that: The response lines of effective events are identified and their positions are encoded. The process includes: within the set field of view angle area, all response line LORs passing through the area are encoded, and each response line LOR has one and only one code corresponding to it. The horizontal axis coordinate of the section of the annular detector imaging device is x, and the vertical axis coordinate of the annular detector imaging device is y. The x and y axes are used to encode the position of each element of the SiPM array respectively. The specific positions of the x and y axes where the annihilation effect occurs in the SiPM array are located respectively through the time window conformity judgment, and the positions are encoded, and the position codes of the two PET detector SiPM arrays are determined, that is, a response line LOR where the annihilation effect occurs is determined.

7. An application of a PET detector imaging system, characterized in that: The image reconstruction is performed by the PET detector imaging system according to claim 1, specifically comprising the following steps: The scintillation crystal LYSO array interacts with the gamma photons and releases the energy in the form of visible light photons; The silicon photomultiplier tube SiPM array converts the visible light signal collected in the corresponding crystal into an electrical signal; The multi-channel PET circuit array converts the electrical signals of the corresponding SiPM array into energy and time signals; The FPGA main control board calculates the location where the effect occurs from the time signal input by the PET circuit array, digitizes the energy signal corresponding to the location and performs Cameralink protocol conversion; The host computer image acquisition card collects the location and energy information of the effect, transmits energy according to the frame rate set by the FPGA, and reconstructs the image.

8. The use of a PET detector imaging system according to claim 7, characterized in that: The FPGA main control board uses an 18-bit 8-channel analog-to-digital conversion chip to digitize the energy analog signal collected by the PET circuit array. The FPGA main control board performs image conversion according to the silicon photomultiplier tube SiPM array, and generates a companion clock CLK, a frame signal FRAME, a row signal ROW, a row valid DAT, and image data respectively. When the row pixel image data of a row starts to be transmitted, the FPGA main control board will generate a high-level row signal ROW and a row valid DAT, and when the row data is completely transmitted, a low-level row signal ROW and a row valid DAT are generated; when the last row of pixel image data of the silicon photomultiplier tube SiPM array is transmitted, the row signal ROW and the row valid DAT are pulled low; when the first row of pixel image data of the first row starts to be transmitted, the FPGA main control board will generate a high-level frame signal FRAME, until all the pixel image data of the silicon photomultiplier tube SiPM array are transmitted, and then the frame signal FRAME is pulled low; the FPGA main control board generates a companion clock CLK every time a pixel data is transmitted until the image is completely transmitted.

9. Use of a PET detector imaging system according to claim 7 or 8, characterized in that: The silicon photomultiplier tube SiPM array is an 8×8 array. Each row of the array is encoded as a group of row pixels. Each row pixel is digitized using 18 bits, and the 18-bit image data of the same row pixel is transmitted in parallel.

10. The use of a PET detector imaging system according to claim 7, characterized in that: When transmitting energy data, the rising edge of the accompanying clock CLK signal is used to collect data. In order to obtain the data output of M×N energy image resolution and frame rate F, it is necessary to meet the following requirements: The silicon photomultiplier tube SiPM array has M rows and N columns; A is the number of valid pixel clock cycles in a row, B is the number of invalid clock cycles in a row, and C is the number of invalid clock cycles in a frame; F clk is the frequency of the accompanying clock CLK signal.

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