The technical solutions of the present invention will be further described below in conjunction with the drawings and specific implementations. It should be understood that in the description of this application, the terms "center", "length", "depth", "thickness", "upper", "lower", "front", "rear", "left", " The orientation or positional relationship indicated by “right”, “vertical”, “horizontal”, “top”, “bottom”, “inner”, “outer”, “clockwise”, “counterclockwise”, etc. are based on the attached drawings The orientation or positional relationship of is only for the convenience of describing the application and simplifying the description, rather than indicating or implying that the device or element referred to must have a specific orientation, be constructed and operated in a specific orientation, and therefore cannot be understood as a limitation of the application .
 The present invention provides a PET detector of a composite crystal array. The structure of the PET detector of the composite crystal array includes: a scintillation crystal module, a photoelectric conversion device module and a detector electronics 11 module. The scintillation crystal module is used to The gamma photons are converted into visible light and soft ultraviolet photons and transmitted to the photoelectric conversion device module. The scintillation crystal module is composed of an A crystal module and a B crystal module to form an AB type mixed light and heavy rare earth single crystal array; the A crystal module is a light rare earth single crystal , B crystal module is a heavy rare earth single crystal; and A and B use a partition structure to form a crystal array. A crystal module is A light rare earth single crystal 3, and B crystal module is B heavy rare earth single crystal 4; the A crystal elements are lanthanum La, cerium Ce, praseodymium Pr, neodymium Nd, promethium Pm, and samarium Sm among the lanthanides. , Europium Eu, gadolinium Gd; the B crystal element is terbium Tb, dysprosium Dy, holmium Ho, erbium Er, thulium Tm, ytterbium Yb, lutetium Lu, scandium Sc, yttrium Y 9 kinds of elements Any kind.
 Rare earth crystals refer to crystals in which rare earth elements can completely occupy a certain lattice point in the crystallographic structure. As a core working substance, they are widely used in laser technology and ionizing radiation detection technology. The mixed light and heavy rare earth single crystal in the present invention means that the crystal array does not use only heavy rare earth elements in the traditional sense, but a mixed use of light and heavy rare earths. The light rare earth refers to lanthanum La and cerium Ce in the lanthanide series. , Praseodymium Pr, neodymium Nd, promethium Pm, samarium Sm, europium Eu, gadolinium Gd 8 kinds of elements, heavy rare earth refers to the lanthanide element terbium Tb, dysprosium Dy, holmium Ho, erbium Er, thulium Tm, ytterbium Yb, lutetium Lu7 elements and two elements closely related to the lanthanides, scandium Sc and yttrium Y.
 The surfaces of the AB two crystal modules have different processing methods. Because when the gamma photons deposit energy inside the crystal to generate scintillation photons, the scintillation photons will propagate inside the crystal and be reflected, refracted or lost on the crystal surface. Part of the photons will eventually be detected by the SiPM detection unit array coupled to the bottom of the crystal. The positioning of the position of the gamma photon and the crystal is mainly realized by the distribution of the output signal of the bottom SiPM detection array. The crystal surface treatment directly determines the behavior of the scintillating photons on the crystal surface and ultimately affects the distribution of the output signal of the SiPM detection array. Therefore, the crystal surface treatment has a vital influence on the three-dimensional position positioning performance of the detector module. Different scintillation photon transport processes may also result in different light collection efficiency of the detector module, so the crystal surface treatment method may also affect the time resolution and energy resolution performance of the detector module. In summary, a proper crystal surface treatment method may play an important role in improving the performance of the continuous crystal detector module. The crystal surface treatment methods mainly include how to treat the crystal surface and what type of reflective film is used to wrap the crystal. Commonly used crystal surface treatments mainly include polishing and emery grinding, while commonly used crystal reflective films mainly include ESR reflective films and Teflon reflective films. In this way, the present invention not only reduces the system cost, but also improves the time resolution and the space resolution, and provides a new idea for the balanced utilization of rare earths.
 The internal basic principle of the PET detector of the composite crystal array is: the radionuclide decays to generate positrons, and the positrons combine with the negative electrons to generate annihilation radiation, generating two gamma photons with 511 keV and opposite flight directions. These two photons have very important properties: they are produced simultaneously in time and fly out almost in opposite directions, which makes it possible to use two oppositely placed detectors outside the body, and use coincidence technology to detect them. Nowadays, they are often closed. The multi-ring detector performs coincidence measurement on these photons. If within the specified time window (usually 0ns~15ns), when the detector detects two photons that are 180 degrees (± 0.25 degrees) of each other, it is a coincidence event, and the annihilation point is the two crystals that flashed. On the line between the blocks, a LOR (Line Of Reaction) response line is formed, and the conformity is recorded in the memory. Since the paths of the two photons in the body are different, the time to reach the respective detectors is also different. From this, the position of the annihilation point on the coincidence line can be calculated. The coincidence circuit system performs coincidence measurement to generate raw data, and completes data acquisition, system monitoring, attenuation correction, and image reconstruction through the computer system.
 According to an embodiment of the present invention, such as figure 1 As shown, a construction method of a composite crystal array PET detector and its construction method is the first scheme, that is, the embedded scintillation crystal array, which includes A crystal and B crystal. The A crystal is a crystal made of cerium-doped light rare earth single crystal lanthanum bromide, and the B crystal is a crystal made of a heavy rare earth single crystal yttrium lutetium silicate. A layer of B crystal is embedded in any two adjacent A crystals, and the number of crystals is increased by analogy to form an n×n crystal array that is an embedded scintillation crystal array. In this example, both the A crystal and the B crystal are It is made of crystals with a square cross-section and has the same shape and size.
 According to an embodiment of the present invention, such as figure 2 As shown, a composite crystal array PET detector and its construction method in the second construction method, that is, the surrounding crystal array 2, including A crystal and B crystal. The A crystal is a crystal made of cerium-doped light rare earth single crystal lanthanum bromide, and the B crystal is a crystal made of a heavy rare earth single crystal yttrium lutetium silicate. Use eight A crystals to enclose one B crystal in the middle to form a 3×3 jiugong lattice crystal array, and by analogy increase the number of jiugong lattices to form an n×n crystal array, which is the surrounding scintillation crystal array. In the example, both the A crystal and the B crystal are made of square-section crystals, and the shapes and sizes are the same.
 According to an embodiment of the present invention, such as image 3 As shown, the length and width of the cerium-doped light rare earth single crystal lanthanum bromide A crystal are 1mm, the height is 3mm, and the embedded heavy rare earth single crystal yttrium lutetium silicate B crystal is 1mm long, 0.2mm wide, and high 3mm.
 According to an embodiment of the present invention, such as Figure 4 As shown, the length and width of the cerium-doped light rare earth single crystal lanthanum bromide A crystal are 1 mm, and the height is 3 mm, and the embedded heavy rare earth single crystal yttrium lutetium silicate B crystal is 1 mm in length and width, and the height is 3 mm.
 According to an embodiment of the present invention, such as Figure 5 As shown, the embedded scintillation crystal array is a 6×6 array with a layer of packaging on the outer surface of the array, and the embedded crystal array package 7 is composed of a reflective film, a coupling layer and a SiPM detection unit. The appearance shape after packaging is cubic.
 According to an embodiment of the present invention, such as Image 6 As shown, the surrounding scintillation crystal array is a 6×6 array, with a layer of packaging on the outer surface of the array, and the surrounding crystal array package 8 is composed of a reflective film, a coupling layer and a SiPM detection unit. The appearance shape after packaging is cubic.
 According to an embodiment of the present invention, Figure 7 The surface of the middle crystal is covered with a reflective film. The reflective film is made of Teflon and aluminum foil to prevent the scintillation photons from being refracted and reflected in the scintillation crystal. There is a coupling layer close to the reflection film. The function of is similar to the light guide, which prevents photons from being too large in the air to directly enter the photomultiplier tube 10. There is a layer of SiPM detection unit at the bottom of the crystal, which is the front end of the photomultiplier tube 10.
 When the PET detector of the composite crystal array is encapsulated, the top, bottom, left, and right surfaces should be wrapped with reflective materials, such as Teflon, BaSO4, etc. The purpose of this is because the scintillation photons will be refracted and reflected in the scintillation crystal. Some photons may also be lost after leaving the scintillator and entering the light guide or the silicone grease layer. The difference in optical length will affect the time characteristics of the scintillation pulse. The reduction in the number of photons in the transport process will affect the energy characteristics of the scintillation pulse. Especially the processing method of the crystal will have an important impact on the reflection and refraction of the scintillating photons. If the surface of the crystal is wrapped with reflective material, part of the emitted photons will be reflected back to the crystal and continue to be transmitted in the crystal.
 According to an embodiment of the present invention, such as Figure 8 As shown, the PET detector of the composite crystal array disclosed in this embodiment includes: a scintillation crystal module, which includes a plurality of crystal modules, a photomultiplier tube module, and a detector electronics module, wherein the detector electronics module includes a first module Electronics and the second module electronics.
 Further, the scintillation crystal module includes an A crystal module and a B crystal module. Both modules are used to absorb and convert the gamma photons emitted from the living body into visible light and soft ultraviolet light photons and transmit them to the photomultiplier tube module. Photocathode module;
 Further, the photomultiplier tube module includes a photocathode module, a focusing electrode module, a dynode module, and a photoanode module. The photocathode module is used to convert the photons transmitted from the A crystal module and the B crystal module into photoelectrons through the photoelectric effect. Pass to the focusing electrode module; the focusing electrode module is used to focus the photoelectron beam and then pass it to the dynode module. The dynode module emits the photoelectron beam twice and multiplies it to the photoanode module. The photoanode module converts the amplified The electrons are collected by the anode as a signal and output to the readout electronics module. Further, the readout electronics module includes a first electronics module and a second electronics module, and the first electronics module includes an amplifier module and an analog-to-digital conversion module, and the second electronics module includes an MVT module, an FPGA module and an Ethernet module. Module. The amplifier module and MVT module both receive electrical signals from the photoanode module, and then the amplifier module amplifies the electrical signals and transmits them to the analog-to-digital conversion module. The analog-to-digital conversion module converts the amplified analog electrical signals into digital electrical signals. The signal is then transmitted to the FPGA module. The MVT module sparsely quantizes the electrical signal and then transmits it to the FPGA module. The FPGA module obtains the signal and then processes it and transmits it to the Ethernet module. After the Ethernet module obtains the signal, it processes and transmits the processed electrical signal. For the data preprocessing module in the image reconstruction and imaging module; the image reconstruction and imaging module 12 includes a data preprocessing module, an image reconstruction module, and an image post-processing and display module. The digital signal obtained from the Ethernet module is transmitted to the data preprocessing module. Module, the data preprocessing module preprocesses the signal and transmits it to the image reconstruction module, and the image reconstruction module transmits the processed signal to the image post-processing and display module, and finally the reconstructed image is obtained.
 Such as Picture 9 As shown, a single crystal array is put into the limit structure designed in advance in the detector ring 9, and the size of the limit structure is equal to the single detector cube module. The detector is the main part of the entire positron emission imaging system, and the detector is composed of a scintillation crystal, a photomultiplier tube 10 and a detector electronics 11. The scintillation crystal and photomultiplier tube 10 are fixed on the detector ring 9. The crystal on the detector ring 9 converts high-energy photons into visible light. The photomultiplier tube 10 connected to the scintillation crystal converts the optical signal into an electrical signal, and the electrical signal It is converted into a time pulse signal, and the compliance circuit checks and judges the time coupling of each crystal pulse signal, eliminates the interference of other sources of rays, and calculates the position of the positron. Each crystal on the closed ring detector has a coincidence relationship with a set of crystals on the opposite side, forming a set of coincidence lines of fan beams. The intersection of all fan beams determines the radial field of view (FOV) of the detector. The two photons generated by the annihilation radiation occurring in the field of view cannot "simultaneously" (within the same time window) hit the same crystal. This type of collimator does not need to be shielded but relies on the special direction and coincidence circuit of the two photons. The method of achieving collimation is called "electronic collimation". PET uses coincidence detection technology for electronic collimation, which greatly reduces random coincidence events and background, and the electronic collimator has very high sensitivity (without the influence of lead shielding) and resolution, which can only be detected simultaneously within a specified time window. Only the two photons captured by the 180° crystals can become a coincidence event. The coincidence circuit confirms that the "coincidence events" entering the same time window are from an annihilation. These "coincidence events" are stored on the specified projection surface, and then the computer uses techniques such as scattering, accidental coincidence signal correction and photon flight time calculation to complete image reconstruction and reconstruction. The image shows the distribution of the tracer in the human body.
 The corners of the A crystal module and the B crystal module are all right angles; the detector package composed of the A crystal module and the B crystal module is a rectangular detector; when the PET detector of the composite crystal array is packaged, the top five The surface should be wrapped with reflective material on the surface. Eight encapsulated PET detectors with self-locking structure are respectively fixed on the detector ring 9 and make the detector modules easy to be inserted in the radial direction.
 A method for constructing a PET detector of the aforementioned composite crystal array, the method includes the following signal transmission steps: step S1: gamma photons are injected into the composite crystal array, and the crystal array absorbs and deposits energy and converts it into visible light; step S2: Visible light enters the photoelectric conversion module, which is converted into an analog electrical signal by the photomultiplier tube; Step S3: The analog electrical signal enters the detector electronics module, and the detector electronics convert it into a digital electrical signal and obtain the time information and Energy information; Step S4: Use computer software to perform data processing and image reconstruction on the obtained time information and energy information, and finally obtain a reconstructed image.
 The keV gamma photons produced by positron annihilation hit the scintillation crystal in the scintillation detector. The gamma photons deposit energy in the scintillation crystal, and then they are converted into many smaller energy scintillation photons. Among them, the scintillation crystal is a functional material that can effectively absorb high-energy rays (X-rays, gamma rays) or high-energy particles and emit ultraviolet and visible light. According to two embodiments of the present invention, the crystals used in the AB two crystal modules are yttrium lutetium silicate LYSO and cerium-doped lanthanum bromide LaBr 3 The combination of these two crystals is a new type of crystal, of which yttrium lutetium silicate LYSO belongs to the monoclinic crystal system, and the space group is C 2 /c, the melting point is as high as 2050℃, has excellent temperature characteristics and physical and chemical properties, high resolution, and faster response; cerium-doped lanthanum bromide LaBr 3 It belongs to the hexagonal crystal system and has obvious anisotropy. It has the advantages of high light output, fast decay time, high energy resolution, and wide application. It is currently one of the best commercial inorganic scintillators.
 The PET detector of a composite crystal array and its construction method are aimed at the current extremely unbalanced rare earth elements and the high-end medical equipment PET uses a large amount of heavy rare earth element lutetium, but the raw material lutetium reserves are small and the price is high. It is difficult to reduce the cost of high-performance PET. This kind of partition design of mixed light and heavy rare earth single crystals is feasible. On the one hand, the afterglow constants of light and heavy rare earth single crystals are different, and higher spatial resolution can be obtained by introducing pulse shape information; on the other hand, light rare earths replace heavy rare earths. Rare earth absorbs radiation, which can reduce system cost. Under the verification of PET design simulation software GATE, the partition plan one, namely figure 1 The embedded crystal array 1 shown has an increase in spatial resolution of 31.2%, while the cost calculated based on the latest rare earth market transaction price has been reduced by 40.1%. In addition, due to the application of cerium-doped light rare earth single crystal LaBr3, there are 29.3% gamma events, and the time resolution is optimized by 2.12 times; with 50.2% gamma events, the time resolution is optimized by 1.53 times. Partition plan two, namely figure 2 In the enclosed crystal array shown, the spatial resolution has increased by 44.2%, and the cost calculated based on the latest rare earth market transaction price has been reduced by 10.3%. Similar to Scheme 1, there are 17.3% of gamma events, and the time resolution is optimized by 2.19 times, and with 31.9% of gamma events, the time resolution is optimized by 1.43 times. As a preliminary exploration, the detector design method provides a new idea for the balanced utilization of rare earths in PET.
 The present invention is described through preferred embodiments, and those skilled in the art know that various changes or equivalent substitutions can be made to these features and embodiments without departing from the spirit and scope of the present invention. The present invention is not limited by the specific embodiments disclosed herein, and other embodiments falling within the claims of the present application all fall within the protection scope of the present invention.