A dual-ended readout pet detector based on continuous crystal light sharing

Abstract

The application discloses a double-end readout PET detector based on continuous crystal light sharing, which comprises N pieces of continuous scintillation crystals, photoelectric converters and signal processing units, and N is greater than or equal to 2; the continuous scintillation crystals are used for blocking gamma rays, converting the gamma rays into scintillation photons and transmitting the scintillation photons to the photoelectric converters; the N pieces of continuous scintillation crystals are arranged into several columns to form an integral structure, and the side surfaces of adjacent continuous scintillation crystals are connected by optical glue; the two end surfaces of each continuous scintillation crystal are connected with a photoelectric converter through optical silicon grease respectively, and are used for generating electrons according to the received scintillation photons and transmitting the electrons to the signal processing units for processing. The application solves the edge effect, thereby being favorable to improving the proportion of the high-resolution area of the detector, improving the positioning precision, having a positive influence on the application of the detector and greatly improving the performance of the PET system.

Description

Technical Field

[0001] This invention relates to the field of nuclear radiation detectors and positron emission tomography (PET) imaging, and in particular to a dual-end readout PET detector based on continuous crystal light sharing and its fabrication method. Background Technology

[0002] Positron emission tomography (PET) is a novel medical instrument invented in the 1970s and widely used in this century. It integrates technologies from physics, mathematics, medicine, electronics, biochemistry, and engineering, playing a crucial role in early tumor diagnosis and examination of neurological diseases, as well as in preclinical animal research. It was named "Invention of the Year" by Time magazine in 2000.

[0003] The imaging principle of PET is: injecting the patient with a radioactive isotope (commonly used) 18 F) of the drug, 18 F decays in a relatively short time. 18 O, simultaneously producing positrons and neutrinos. The decay formula is:

[0004] p→n+e + +v+energy#(1.1.1)

[0005] Positrons have the same mass as electrons but opposite charges, making them a type of antimatter. In tissues, positrons continuously undergo elastic collisions with outer electrons, losing kinetic energy and eventually annihilating each other to produce two gamma photons. These two gamma photons each have an energy of 511 keV and are oriented approximately 180° apart. The detector detects the locations where these two gamma photons interact within a coincidence time window; the line connecting these two locations is called the line of view (LOR). The intersections of numerous LORs indicate the locations where positron annihilation occurs. The PET system reconstructs these positron annihilation locations using image reconstruction algorithms, thereby acquiring image information.

[0006] Currently, gamma ray detectors usually adopt the configuration of coupling crystal array with position sensitive photoelectric device, combined with Anger logic algorithm to realize the positioning inside the detector. The pixel size of the crystal array corresponds to the planar positioning accuracy, the smaller the size, the stronger the positioning ability. In current various types of nuclear medical imaging instruments, the position resolution of the gamma ray detector is mostly above mm. With the continuous development of PET technology and the demand for precision medicine and dynamic imaging, high system spatial resolution and high system sensitivity have become the mainstream of development. Currently, the pixel size of the detector used in clinical human PET systems is mostly 2.5-4 mm; in preclinical small animal PET, the pixel size is mostly 1.0-2.0 mm. Especially in preclinical animal PET systems, people are pursuing higher and higher spatial resolution and sensitivity. The high sensitivity and high spatial resolution of the system require high planar position resolution and high detector efficiency of the detector.

[0007] The high planar position resolution of the detector corresponds to the small size of the crystal unit. With the continuous progress of detector technology, the position decoding of a 0.32 mm unit LYSO crystal array has been realized. However, due to the existence of about 0.1 mm optical reflection layer between the crystal units, the detection efficiency is lost. For a 0.5 mm unit size crystal array, the efficiency loss of a single detector reaches 30% under the same detector area, and the coincidence efficiency loss reaches 51%.

[0008] In preclinical PET systems, due to the small ring diameter, in addition to the requirement of high planar position resolution of the detector, DOI resolution capability is also required, i.e. three-dimensional position detector. When the gamma ray generated by positron annihilation obliquely enters the detector, if the detector does not have DOI resolution capability, only the top surface of the crystal unit can be used as the starting point and the ending point of the response line, causing the reconstruction of the annihilation position to be offset, i.e. parallax error. The parallax error phenomenon becomes serious with the reduction of the PET ring diameter, so the DOI resolution capability is most required in small ring diameter PET. Currently, most commercial preclinical PET systems do not have DOI resolution capability. When not having DOI resolution capability, the image will become blurred, and the spatial resolution of the system gradually deteriorates from the center to the edge of the FOV.

[0009] The continuous crystal is a whole piece without cutting, and there is no reflection layer between the crystal units, so there is no problem of detection efficiency loss, and at the same time, a very high three-dimensional position resolution can be achieved under certain detector parameters. In addition, the upper limit of the three-dimensional spatial positioning capability of the continuous crystal detector is high, and when a more reasonable detector configuration and positioning algorithm is adopted, the three-dimensional position resolution will exceed that of other types of detectors, so it is easy to achieve high detection efficiency and high three-dimensional position resolution. Compared with the multi-layer detector stack, the electronic structure used by the continuous crystal detector is simpler, and there is no scattering caused by a large number of readout layers. The DOI detector based on the stacking of crystal arrays is mostly limited by the process and discrimination capability, resulting in a limited number of layers and low DOI resolution. In the light sharing between array crystal units and the top light sharing scheme, it is also limited by its own structure, and it is difficult to achieve high DOI resolution at high planar position resolution. In addition to high detection efficiency and high three-dimensional position resolution, the continuous crystal detector also has excellent time-of-flight (TOF) potential. Since the scintillation photons do not have to bear the influence of internal reflection in the crystal array unit, the first photon can directly reach the photodetector, so it has better time performance compared with the crystal array type detector.

[0010] As described above, the continuous crystal detector has been applied to a certain extent in small animal PET systems. How to further improve the three-dimensional positioning accuracy and detection efficiency of the continuous crystal detector has become the focus. For improving the detection efficiency, the simplest way is to increase the thickness of the scintillation crystal. According to the Beer-Lambert law, the detection efficiency of a 20mm thick LYSO crystal for 511keV gamma rays is 81%. Although a 20mm thick crystal is very common in clinical PET, its application in preclinical animal PET devices is still less. Therefore, using a thicker continuous crystal can improve the sensitivity of preclinical animal PET. However, when the thickness increases, the transmission range of the photons in the crystal is wider, the distribution on the SiPM array plane is wider, and even it can exceed the SiPM plane, and the edge effect is aggravated. In another aspect, when the light distribution is wider, the number of photons received by each SiPM is less, the signal-to-noise ratio is reduced, which is not conducive to energy resolution and three-dimensional position resolution. Taking the case where a gamma ray deposits energy at a distance of 20mm from the SiPM plane, the SiPM directly below can only detect 10-20 photons. Since most gamma rays interact in the first half of the crystal, the performance of the near-end readout is superior compared to the traditional SiPM far-end readout.

[0011] Further, the double-end readout scheme that SiPM array readout is coupled at both ends of the crystal can achieve better three-dimensional position resolution in thick crystals. Studies have shown that using PDPC to couple to a 22m thick LYSO crystal and using the k nearest neighbor algorithm, the single-end readout detector can obtain a 1.7mm plane resolution and a 3.7mm DOI resolution, while the double-end readout can improve the resolution to 1.1mm and 2.4mm respectively. It can be seen that for the purpose of achieving high three-dimensional positioning accuracy and high detection efficiency of the PET detector, the double-end readout scheme based on continuous crystals is a very feasible scheme.

[0012] Because the distribution of scintillation light in the continuous crystal is truncated at the edge of the crystal, the energy resolution, three-dimensional position resolution and other performances at this position are poor. SUMMARY

[0013] In view of the problems existing in the existing PET detector, combined with the analysis of the advantages of the continuous crystal detector applied to the PET system, the purpose of the present application is to provide a double-end readout PET detector based on continuous crystal light sharing, which solves the edge effect, thereby improving the proportion of high resolution area of the detector and improving the positioning accuracy, and has a positive influence on the application of the detector. The detector has the characteristics of high three-dimensional spatial resolution and high detection efficiency. Based on such a detector module, a PET system with high spatial resolution, DOI resolution capability and ultra-high sensitivity can be efficiently and quickly built, and the system has higher 2D resolution in the axial direction, greatly improving the performance of the PET system.

[0014] The technical scheme of the present application is:

[0015] A double-end readout PET detector based on continuous crystal light sharing, characterized in that it comprises N continuous scintillation crystals, a photoelectric converter and a signal processing unit, and N is greater than or equal to 2.

[0016] The continuous scintillation crystal is used to block gamma rays, convert the gamma rays into scintillation photons and transmit them to the photoelectric converter.

[0017] The N continuous scintillation crystals are arranged in several columns to form an overall structure, and the sides of adjacent continuous scintillation crystals are connected by optical glue.

[0018] The two end surfaces of each continuous scintillation crystal are connected to a photoelectric converter through optical silicon grease, respectively, for generating electrons according to the received scintillation photons and transmitting them to the readout electronics circuit.

[0019] The signal processing unit is used to process and calculate the signals of the photoelectric converter.

[0020] Further, the side surface of the whole structure is coated with a black pigment coating for absorbing the scintillation photons reaching the side boundary of the whole structure.

[0021] Further, the thickness of the black pigment coating is 0.25 mm.

[0022] Further, the photoelectric converter is a silicon photomultiplier array, each of which is used for receiving the scintillation photons reaching one end of the continuous scintillation crystal.

[0023] Further, the photoelectric conversion device is a photomultiplier tube.

[0024] Further, the continuous scintillation crystal is a continuous LYSO crystal.

[0025] Further, the refractive index of the optical glue is 1.70.

[0026] Further, the signal processing unit includes readout electronics circuit, analog electronics and digital electronics; or the signal processing unit is an ASIC readout chip.

[0027] A positioning method based on the double-end readout PET probe, the steps of which include:

[0028] 1) Arrange N pieces of the continuous scintillation crystal along the X direction into a column to form a whole structure; the photoelectric converter is a silicon photomultiplier array composed of M×M silicon photomultiplier units, each of which is used for receiving the scintillation photons reaching one end of the continuous scintillation crystal; the nth silicon photomultiplier array is used for receiving the scintillation photons reaching the nth continuous scintillation crystal Block n far end, the 2nth silicon photomultiplier array is used for receiving the scintillation photons reaching the nth continuous scintillation crystal Block n near end, n=1~N; define the nth continuous scintillation crystal Block n middle far end as nF end, the nth continuous scintillation crystal Block n middle near end as nf end, the nth silicon photomultiplier array receives the signal in the X direction at the nF end of Block n . nFX1 , nFX2 , nFXM ……, nFY1 , nFY2 , nFYM ; the 2nth silicon photomultiplier array receives the signal in the X direction at the nf end of Block n . nfX1 , nfX2 , nfXM ……,The signal received in the Y direction is denoted as E. nfY1 E nfY2 ... E nfYM ;

[0029] 2) If the scintillation light is generated at both ends of the overall structure, it is a continuous scintillation crystal block. j j = 1 or N, that is, the first continuous scintillation crystal Block1, the Nth continuous scintillation crystal Block N Block j The distal and proximal ends are denoted as jF end and jf end, respectively, and are related to Block. j Adjacent Continuous Scintillation Crystal Block j’ The distal and proximal ends are denoted as j'F end and j'f end, respectively; then for Block j The following results are obtained by subtracting the threshold from each signal in the X direction at the jF end:

[0030]

[0031] And Block j’ The following results are obtained by subtracting the threshold from each signal in the X direction at the j'F end:

[0032] β is a set threshold;

[0033] 3) Follow the steps in step 2) to modify the Block. j The signals in the Y direction at the jF end are subjected to threshold reduction processing, and the Block is processed. j The X-direction and Y-direction signals at the jf terminal are subjected to threshold reduction processing;

[0034] 4) For the X direction: using the Block obtained in steps 2) and 3) j and its adjacent continuous scintillation crystal Block j’ After threshold reduction, the far-end X-direction signal is used to calculate the far-end positioning result in the X-direction using the centroid method, denoted as X. jF Using step 2)

[0035] 3) The obtained Block j and its adjacent continuous scintillation crystal Block j’ After threshold reduction, the X-direction signal of the proximal end is used to calculate the proximal end localization result in the X direction using the centroid method, denoted as X. jf Then calculate the Block j X-direction localization result of internal scintillation light j =(X jF +X jf ) / 2; similarly, the Block is calculated.j Y-direction localization result of internal scintillation light j Thus, we obtain the Block. j Location of internal scintillation light generation (X) j Y j );

[0036] 5) If the scintillation light is generated in a continuous scintillation crystal block within the overall structure k Let the distal and proximal ends be denoted as kF end and kf end, respectively, where k = 2 to N-1; then, according to the Block... k Adjacent Continuous Scintillation Crystal Block k’ Block k” Signals in the X direction affect Block k The threshold is subtracted from the kF terminal X-direction signal to obtain:

[0037]

[0038] 6) Follow the steps in step 5) to process the Block k The kF terminal Y-direction signals are subjected to threshold reduction processing, and the Block signal is processed. k The thresholds of each signal in the X direction and each signal in the Y direction at the kf terminal are reduced.

[0039] 7) For the X direction: using the Block obtained in steps 5) and 6) k And its adjacent continuous scintillation crystal Block k’ and Block k” The remote X-direction signal after threshold reduction is used to calculate the remote X-direction positioning result using the centroid method, denoted as X. kF Using the Block obtained in steps 5) and 6) k and its adjacent continuous scintillation crystal Block k’ and Block k” The X-direction localization result is obtained by calculating the proximal X-direction signal after threshold reduction using the centroid method, and is denoted as X. kf Then calculate the Block k X-direction positioning results k =(X kF +X kf ) / 2; similarly, the Block is calculated. k Y-direction localization result of internal scintillation light k Thus, we obtain the Block. k Location of internal scintillation light generation (X) k Y k ).

[0040] A PET system characterized in comprising a plurality of the above-mentioned dual-ended readout PET detectors.

[0041] The present application is based on a dual-ended readout PET detector module with continuous crystal light sharing, which comprises N pieces of continuous scintillation crystals (N≥2), a photoelectric converter, and a signal processing unit (traditional electronic circuit: including readout electronics, analog electronics, and digital electronics; or an ASIC readout chip), which together constitute a detector module; the detector module can be combined in any number (≥2) to form a PET system, and the detector modules are connected by optical fibers. Figure 1 The planar structure of the detector is shown in

[0042] In an embodiment of the present application, the continuous scintillation crystals are used to block gamma rays, convert the gamma rays into scintillation photons, and transmit them to the photoelectric conversion device.

[0043] Specifically, the three-dimensional size of the continuous scintillation crystals is not required, and the size of the crystals is arbitrary. Further, in an embodiment of the present application, continuous LYSO crystals are used, and the side surfaces are sequentially connected between the N pieces of continuous scintillation crystals using optical glue with high refractive index, so that the N pieces of continuous scintillation crystals are spliced into a whole. The present application does not limit the arrangement mode of the N pieces of continuous scintillation crystals, and according to the requirements of the user for the shape and size of the detector, the N pieces of continuous scintillation crystals can be arranged in a row or in multiple rows.

[0044] Further, for the N pieces of continuous scintillation crystals after connection, all the exposed side surfaces (i.e. the side surfaces not subjected to bonding) are uniformly painted black with black pigment.

[0045] The structure of the continuous scintillation crystals after splicing is shown in Figure 2

[0046] In an embodiment of the present application, the photoelectric conversion device is a silicon photomultiplier array or a photomultiplier tube; if it is a silicon photomultiplier array, a total of 2N silicon photomultiplier arrays are required, each silicon photomultiplier array is composed of M 2 silicon photomultiplier units with an area of (1~S)×(1~S) mm 2 , and is used to convert scintillation photons into photoelectrons and output pulse electrical signals after multiplication of the photoelectrons.

[0047] Each silicon photomultiplier unit corresponds to one energy signal, and each silicon photomultiplier array corresponds to M 2 energy signals.

[0048] The time signals of each silicon photomultiplier unit are summarized into one, and each silicon photomultiplier array corresponds to one time signal.

[0049] ​Further, the upper and lower surfaces of the spliced N continuous scintillation crystals are coupled with 2N silicon photomultiplier tube arrays through optical silicon grease to achieve a double-end readout effect.

[0050] In an embodiment of the application, the signal processing unit is a conventional electronic circuit, including readout electronics, analog electronics, and digital electronics.

[0051] The readout electronics is a symmetrical charge division circuit, which is a resistance network structure that converts M 2 channel energy signals into 2M channel outputs.

[0052] The analog electronics includes an amplifier module, a high-voltage power supply module, a low-voltage stabilizing module, and a fast signal processing module, which are used to power the front-end silicon photomultiplier tube array, amplify the 2M channel energy signals, and perform front edge timing or constant ratio timing on the 1 channel time signal.

[0053] The digital electronics is used to integrate the amplified energy waveform to obtain the energy signal E i , and sample the timed time signal as a trigger signal, and finally transmit all the energy and time information to a computer. Figure 3

[0054] A PET system structure composed of the detector of the application is shown in Figure 4 .

[0055] To achieve the purpose of the application, the application further provides a calculation method for fine three-dimensional positioning of a continuous crystal detector, which takes a detector composed of four continuous scintillation crystals bonded and spliced along a method as an example, and in the example, the number of silicon photomultiplier tube arrays used by the detector is MxM. Users can apply this calculation method to detectors of other bonding methods according to the example.

[0056] Step S1: Definition. As shown in Figure 5 , define the X-direction signals of the A end (far end) of the first continuous scintillation crystal Block1 as E AX1 , E AX2 , …, E AXM , the Y-direction signals as E AY1 , E AY2 , …, E AYM , the X-direction signals of the a end (near end) of Block1 as E aX1 , E aX2 , …, E aXM , and the Y-direction signals as E aY1 , E aY2 , …, E aYM ​; second continuous scintillation crystal Block2, third continuous scintillation crystal Block 3, fourth continuous scintillation crystal Block 4 are the same.

[0057] Step S2: Channel threshold reduction.

[0058] For X direction:

[0059] If the scintillation light generation position is in Block1: (Block4 and Block1 are the same, and for Block4, the C end signal of the adjacent Block3 is needed to process the D end signal of Block4);

[0060] After reducing the threshold of each signal of Block1 A end X direction:

[0061] The signal of the adjacent Block in X direction is:

[0062] If the scintillation light generation position is in Block2: (Block3 and Block2 are the same);

[0063] After reducing the threshold of each signal of Block2 B end X direction:

[0064] The signal of the adjacent two Blocks in X direction:

[0065] For Y direction:

[0066] Block1, Block 2, Block 3, Block 4 are logically the same, taking Block1 as an example:

[0067] Wherein, β in all the above formulas is a threshold, which can be any value less than 1.

[0068] Step S3: Calculate the position.

[0069] If the scintillation light is generated in the two-end continuous scintillation crystal Block j , j = 1 or N, i.e. the first continuous scintillation crystal Block1, the Nth continuous scintillation crystal Block N ;

[0070] For X direction: use Block j and its adjacent continuous scintillation crystal Block j’The far-end X-direction signal after threshold reduction is used to perform the centroid method to obtain the far-end localization result in the X direction, denoted as X. jF ; Using Block j And its adjacent continuous scintillation crystal Block j’ The centroid method is used to perform proximal X-direction localization on the subtracted threshold signal. The proximal X-direction localization result is denoted as X. jf ; final Block j The X-direction positioning result is X j =(X jF +X jf ) / 2.

[0071] For the Y direction: the calculation method for the X direction is analogous to obtain the Block. j Y-direction localization result of internal scintillation light j ; ultimately obtained Block j Location of internal scintillation light generation (X) j Y j );

[0072] If the scintillation light is generated in the middle of the continuous scintillation crystal Block k k = 2 to N-1;

[0073] For the X direction: using Block k And its adjacent continuous scintillation crystal Block k’ and Block k” The far-end X-direction signal after threshold reduction is used to perform the centroid method to obtain the far-end localization result in the X direction, denoted as X. kF ; Using Block k And its adjacent continuous scintillation crystal Block k’ and Block k” The proximal X-direction signal after threshold reduction is used to perform centroid method to obtain the proximal X-direction localization result, denoted as X. kf ; final Block k The X-direction positioning result is X k =(X kF +X kf ) / 2.

[0074] For the Y direction: the calculation method for the X direction is analogous to calculate the Block. k Y-direction localization result of internal scintillation light k Ultimately, we obtained the Block. k Location of internal scintillation light generation (X) k Y k ).

[0075] The present invention has the following beneficial effects:

[0076] 1) The application adopts a continuous scintillation crystal detector, compared with a traditional array type crystal detector, reduces the detection efficiency loss caused by the light isolation, so that the application has higher detection efficiency;

[0077] 2) The continuous crystal light sharing scheme used in the application makes the scintillation light pass through the interface between adjacent continuous crystals and be detected by the corresponding photodetector of the adjacent crystal, realizes the partial recovery of the light distribution truncation at the edge, and improves the spatial resolution of the detector at the edge;

[0078] 3) The double-end readout scheme used in the application makes the three-dimensional spatial resolution of the continuous crystal higher, and ensures that a higher spatial resolution can be achieved in the case of a 20mm thick crystal, thereby improving the three-dimensional spatial resolution and detection efficiency of the detector;

[0079] 4) The number of spliced crystals used in the detector described in the application can be selected by the user, the modular design is free to assemble into a PET system with different axial lengths, and the use is more convenient. BRIEF DESCRIPTION OF DRAWINGS

[0080] Figure 1 It is a planar structure diagram of the detector.

[0081] Figure 2 It is a schematic diagram of the crystal splicing structure of the detector.

[0082] Figure 3 It is a whole structure diagram of the detector described in the application.

[0083] Figure 4 It is a structure diagram of a PET system composed of the detector of the application.

[0084] Figure 5 It is a schematic diagram used in a fine three-dimensional spatial positioning calculation method of a double-end readout detector for continuous crystal light sharing. DETAILED DESCRIPTION

[0085] The following description is used to disclose the application so that those skilled in the art can implement the application. The preferred embodiments in the following description are only used as examples, and other obvious modifications can be thought of by those skilled in the art. The basic principles of the application defined in the following description can be applied to other embodiments, modifications, improvements, equivalents and other technical solutions without departing from the spirit and scope of the application.

[0086] Those skilled in the art should understand that in the disclosure of the present application, the terms "vertical", "transverse", "upper", "lower", "front", "back", "left", "right", "vertical", "horizontal", "top", "bottom", "inner", "outer" and the like indicate the orientation or positional relationship based on the orientation or positional relationship shown in the drawings, which is only for the convenience of describing the present application and simplifying the description, and does not indicate or imply that the device or element referred to must have a particular orientation, be constructed and operated in a particular orientation, therefore the above terms cannot be understood as a limitation of the present application.

[0087] It can be understood that the term "one" should be understood as "at least one" or "one or more", that is, in one embodiment, the number of one element can be one, and in another embodiment, the number of the element can be multiple, and the term "one" cannot be understood as a limitation on the number.

[0088] In the description of the present application, it should be noted that unless otherwise explicitly specified and limited, the terms "mounting", "connection", "connection" should be understood broadly, for example, it can be fixedly connected, or it can be detachably connected, or integrally connected; it can be mechanically connected, or it can be electrically connected or can communicate with each other; it can be directly connected, or it can be indirectly connected through an intermediate medium, it can be the internal communication of two elements or the interaction relationship between two elements. For those skilled in the art, the specific meaning of the above terms in the present application can be understood according to the specific circumstances.

[0089] The present application designs a double-end readout PET detector based on continuous crystal light sharing in one aspect. The detector has high three-dimensional spatial resolution and high detection efficiency, and uses any number of the detector to form a PET system, which has higher system spatial resolution, can eliminate the parallax caused by DOI effect, and has higher system sensitivity compared with the traditional PET system. Moreover, the detector adopts modular design, can be arbitrarily composed into PET systems with different ring diameters and axial lengths, can be suitable for various imaging scenes, and is more convenient to use.

[0090] With Figure 1 The typical embodiments shown correspond to the detector, and the specific implementation of the present application is described:

[0091] The detector is composed of four 40mm*40mm*20mm continuous LYSO crystals spliced by side surface, the refractive index of optical glue used for splicing is 1.70; the exposed surface of the spliced crystal is uniformly blackened by black pigment except the upper and lower bottom surface, and the thickness of the black pigment coating is about 0.25mm; the upper and lower bottom surface of the crystal is coupled with 8 12*12 SiPM arrays through optical silicone grease; the whole crystal and SiPM array are put into the aluminum bracket at the front end of the detector, and a layer of rubber buffer material is added between the SiPM array PCB and the bracket; each SiPM array is connected with a readout electronics board, and the readout electronics board is a PCB with flexible wires, so that the signals of the double-end readout are introduced to the rear end.

[0092] It can be understood that in this specific example of the application, four 40mm*40mm*20mm continuous LYSO crystals are used, and the user can customize the number of continuous crystals used in the detector according to actual needs, as long as the number of continuous crystals is greater than or equal to 2, which mainly affects the detection area of the detector and the field of view range when used in a PET system, and the application does not limit this.

[0093] Specifically, the thickness D of the crystal is 1-20mm, preferably 20mm, which is to ensure the high detection efficiency of the detector.

[0094] Specifically, the SiPM array in the application is composed of a plurality of silicon photomultipliers with an area of (1-3) * (1-3) mm 2 , and the thickness of the SiPM array is 1mm.

[0095] That is, the SiPM array is composed of a plurality of unit silicon photomultipliers, and the unit area of the unit silicon photomultiplier can be 1*1mm 2 , 2*2mm 2 , or 3*3mm 2 , and the application does not limit this.

[0096] In particular, in this typical embodiment of the application, the SiPM array can be a 12*12 SiPM array with a unit size of 3*3mm 2 , and the overall size of this SiPM array is 40*40*1mm 3 . This SiPM array has the characteristics of small area and small dead zone, which is beneficial to ensure the detection performance of the detector.

[0097] Specifically, the optical glue used in the present application is used to splice two adjacent continuous crystals into a whole, so that the scintillation photons reaching the boundary of the scintillation crystal can be detected by the adjacent crystal, and the optical glue should ensure high transmission efficiency of scintillation light and similar refractive index to the refractive index of the scintillation crystal. In this specific example of the present application, optical glue with a refractive index of n=1.70 is used, and the user can appropriately adjust the type of optical glue to meet the above requirements, and the present application does not limit this.

[0098] Specifically, the "blackening the exposed surfaces of the spliced crystal except the upper and lower surfaces with black paint" described in the present application is used to absorb the scintillation photons reaching the side boundary. The selected black paint should ensure high absorption rate of photons in the visible light band. The present application does not limit the type and model of the selected black paint.

[0099] Specifically, the upper surface (proximal end) and the lower surface (distal end) of the N continuous scintillation crystals after connection are coupled to the 2N silicon photomultiplier tube array through optical silicone.

[0100] Specifically, the upper and lower surfaces of all scintillation crystals and the surfaces used for connection with other crystals described in the present application are polished, and the polishing of the other surfaces coated with black paint is not limited.

[0101] Another aspect of the present application provides a fine three-dimensional spatial positioning scheme of a double-end readout detector for continuous crystal light sharing:

[0102] According to the performance test results, the double-end readout detector for continuous crystal light sharing of the present application has higher detection efficiency, sub-millimeter planar resolution and less than 2mm depth resolution compared with the traditional array detector, and the performance is obviously improved.

[0103] The present application has high detection efficiency for positron imaging and at the same time ensures high three-dimensional position resolution of the detector, and has high application value in small animal PET systems.

[0104] Although specific embodiments of the present application are disclosed for the purpose of illustration and description, the purpose is to help understand the content of the present application and to implement it, and those skilled in the art can understand that various substitutions, changes and modifications are possible without departing from the spirit and scope of the present application and the appended claims. Therefore, the present application should not be limited to the disclosed content of the best embodiment, and the scope of the present application claimed is defined by the scope of the claims.

Claims

1. A method of localizing a dual-ended readout PET detector, wherein, The dual-end readout PET probe includes N pieces of continuous scintillation crystals for blocking gamma rays, converting the gamma rays into scintillation photons and transmitting the scintillation photons to a photoelectric converter, N≥2; The steps include: 1) N pieces of the continuous scintillation crystal are arranged in a column along the X direction to form an integral structure; the photoelectric converter is a silicon photomultiplier array composed of MxM silicon photomultiplier units, each of which is used to receive scintillation photons reaching one end of a continuous scintillation crystal; the nth silicon photomultiplier array is used to receive scintillation photons reaching the nth continuous scintillation crystal Block n far end, the 2nth silicon photomultiplier array is used to receive scintillation photons reaching the nth continuous scintillation crystal Block n near end, n = 1 ~ N; the nth continuous scintillation crystal is defined as Block n the middle far end is nF end, the nth continuous scintillation crystal is Block n the middle near end is nf end, the nth silicon photomultiplier array receives signals in the X direction at the nF end of Block n , which are denoted as E nFX1 , E nFX2 , …, E nFXM , signals received in the Y direction are denoted as E nFY1 , E nFY2 , …, E nFYM ; the 2nth silicon photomultiplier array receives signals in the X direction at the nf end of Block n , which are denoted as E nfX1 , E nfX2 , …, E nfXM , signals received in the Y direction are denoted as E nfY1 , E nfY2 , …, E nfYM ; 2) If the scintillation light is generated in the continuous scintillation crystal Block j , j = 1 or N, i.e. the first continuous scintillation crystal Block1, the Nth continuous scintillation crystal Block N ; the far end and the near end of Block j are respectively recorded as jFend and jfend, and the far end and the near end of the continuous scintillation crystal Block j adjacent to Block j’ are respectively recorded as j'Fend and j'fend; then the jFend X-direction signals of Block j are obtained by subtracting the threshold value from the signals: and subtracting the threshold from each of the signals in the X direction of the j'F end of Block j’ ​ Beta is a set threshold value; 3) Threshold reduction is performed on each signal of the jF end Y direction of Block j in the same manner as in step 2); and threshold reduction is performed on each signal of the jF end X direction and each signal of the jF end Y direction of Block j . 4) For the X direction: using the Block obtained in steps 2) and 3) j and its adjacent continuous scintillation crystal Block j’ After threshold reduction, the far-end X-direction signal is used to calculate the far-end positioning result in the X-direction using the centroid method, denoted as X. jF Using the Block obtained in steps 2) and 3) j and its adjacent continuous scintillation crystal Block j’ After threshold reduction, the X-direction signal of the proximal end is used to calculate the proximal end localization result in the X direction using the centroid method, denoted as X. jf Then calculate the Block j X-direction localization result of internal scintillation light j =(X jF +X jf ) / 2; similarly, the Block is calculated. j Y-direction localization result of internal scintillation light j Thus, we obtain the Block. j Location of internal scintillation light generation (X) j Y j ); 5) If the scintillation light is generated in the continuous scintillation crystal Block k , the far end, the near end are recorded as kF end, kf end, respectively, k = 2 ~ N-1; then according to the Block k , the far end, the near end are recorded as kF end, kf end, respectively, k = 2 ~ N-1; then according to the Block k’ , the far end, the near end are recorded as kF end, kf end, respectively, k = 2 ~ N-1; then according to the Block k” , the far end, the near end are recorded as kF end, kf end, respectively, k = 2 ~ N-1; then according to the Block k , the far end, the near end are recorded as kF end, kf end, respectively, k = 2 ~ N-1; then according to the Block 6) Threshold reduction is performed on each signal of Block k in the Y direction at the kF end in the same manner as in step 5), and on each signal of Block k in the X direction at the kF end and in the Y direction. 7) For the X direction: using the Block obtained in steps 5) and 6) k And its adjacent continuous scintillation crystal Block k’ and Block k” The remote X-direction signal after threshold reduction is used to calculate the remote X-direction positioning result using the centroid method, denoted as X. kF Using the Block obtained in steps 5) and 6) k and its adjacent continuous scintillation crystal Block k’ and Block k” The X-direction localization result is obtained by calculating the proximal X-direction signal after threshold reduction using the centroid method, and is denoted as X. kf Then calculate the Block k X-direction positioning results k =(X kF +X kf ) / 2; similarly, the Block is calculated. k Y-direction localization result of internal scintillation light k Thus, we obtain the Block. k Location of internal scintillation light generation (X) k Y k ).

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