Double-end reading PET detector based on continuous crystal light sharing

Abstract

The invention discloses a double-end read-out PET detector based on continuous crystal light sharing, which comprises N continuous scintillation crystals, a photoelectric converter and a signal processing unit, and N is greater than or equal to 2; the continuous scintillation crystal is used for blocking gamma rays, converting the gamma rays into scintillation photons and transmitting the scintillation photons to the photoelectric converter; the N continuous scintillation crystals are arranged into a plurality of columns to form an integral structure, and the side surfaces of the adjacent continuous scintillation crystals are connected through optical cement; the two end faces of each continuous scintillation crystal are connected with a photoelectric converter through optical silicone grease, and the continuous scintillation crystals are used for generating electrons according to received scintillation photons and transmitting the electrons to a signal processing unit for processing. According to the method, the edge effect is solved, so that the high-resolution area proportion of the detector is improved, the positioning precision is improved, positive influences are generated on the application of the detector, and the performance of a PET system is greatly improved.

Description

Technical Field

[0001] The present invention relates to the field of nuclear radiation detectors and positron emission tomography (PET) imaging, and in particular to a double-end readout PET detector based on continuous crystal light sharing and a manufacturing method thereof. Background Art

[0002] Positron Emission Tomography (PET) is a new type of medical device invented in the 1970s and widely used in this century. It integrates technologies from physics, mathematics, medicine, electronics, biochemistry and engineering, and plays an important role in early tumor diagnosis and examination of brain and nervous system diseases, as well as in preclinical animal research. It was selected as the "Invention of the Year" by Time magazine in 2000.

[0003] The imaging principle of PET is: injecting a radioactive isotope (usually 18 F) drugs, 18 F will decay in a shorter time into 18 O, producing positrons and neutrinos at the same time. The decay formula is:

[0004] p→n+e + +v+energy#(1.1.1)

[0005] Positrons have the same mass as electrons but opposite charges, and are a type of antimatter. Positrons constantly collide elastically with extranuclear electrons in tissues, lose kinetic energy, and eventually annihilate with electrons, producing two gamma photons. The energy of these two gamma photons is 511keV, and their directions are approximately 180°. The detector detects the position where two gamma photons interact in the detector within the time window. The line between the two positions is called the line of view (LOR), and the intersection of a large number of line of view lines is the position where the positron annihilates. The PET system reconstructs the positron annihilation position through an image reconstruction algorithm to obtain image information.

[0006] At present, gamma-ray detectors usually adopt the configuration of crystal array coupled with position-sensitive photoelectric devices, combined with Anger logic algorithm to achieve positioning inside the detector. The pixel size of the crystal array corresponds to the plane positioning accuracy. The smaller the size, the stronger the positioning ability. In the current various types of nuclear medicine imaging instruments, the position resolution of gamma-ray detectors 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. At present, the pixel size of the detector used in the human PET system used in clinical practice is mostly 2.5mm-4mm; in preclinical small animal PET, the pixel size is mostly 1.0-2.0mm. 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 plane position resolution and high detector efficiency.

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

[0008] In preclinical PET systems, due to the small ring diameter, in addition to requiring the detector to have high planar position resolution, DOI resolution is also required, that is, a detector with three-dimensional position. When the gamma rays generated by positron annihilation are obliquely incident on the detector, if the detector does not have DOI resolution, it can only use the top surface of the crystal unit as the starting and ending point of the response line, causing the reconstruction of the annihilation position to shift, that is, parallax error. The parallax error phenomenon becomes more serious as the PET ring diameter decreases, so DOI resolution is most demanded in small ring diameter PET, but most current commercial preclinical PET systems do not have DOI resolution. When DOI resolution is not available, the image becomes 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. There is no reflection layer between crystal units, so there is no problem of detection efficiency loss. At the same time, it can also achieve high three-dimensional position resolution under specific detector parameters. In addition, the upper limit of the three-dimensional spatial positioning capability of the continuous crystal detector is very high. When a more reasonable detector configuration and positioning algorithm are adopted, the three-dimensional position resolution will surpass other types of detectors, so it is easy to achieve high detection efficiency and high three-dimensional position resolution. Compared with multi-layer detector stacking, the electronic structure used in the continuous crystal detector is simpler, and there is no scattering caused by a large number of readout layers. Most DOI detectors based on crystal array stacking are limited by process and discrimination capabilities, resulting in limited number of layers and low DOI resolution. In the light sharing between array crystal units and the light sharing scheme at the top, it is also limited by its own structure, and it is difficult to achieve high DOI resolution at high plane 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 withstand the reflection inside the crystal array unit, the first photon can directly reach the photodetector, so it has better time performance than the crystal array detector.

[0010] As mentioned above, continuous crystal detectors have been applied to a certain extent in small animal PET systems. How to further improve the three-dimensional positioning accuracy and detection efficiency of continuous crystal detectors has become the focus. For improving 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 20mm thick crystals are very common in clinical PET, they are still rarely used in preclinical animal PET equipment. Therefore, the use of thicker continuous crystals can improve the sensitivity of preclinical animal PET. However, when the thickness increases, the wider the transmission range of photons inside the crystal, the wider the distribution in the SiPM array plane, and even beyond the SiPM plane, aggravating the edge effect. From another perspective, when the light distribution is wider, the number of photons received by each SiPM is fewer, the signal-to-noise ratio is reduced, and it is not conducive to energy resolution and three-dimensional position resolution. Taking the energy deposition of γ at 20mm from the SiPM plane as an example, the SiPM directly below can only detect 10-20 photons. Since most of the gamma rays interact in the front half of the crystal, the performance of proximal readout is superior compared to conventional SiPM distal readout.

[0011] Furthermore, a dual-end readout scheme that couples the SiPM array readout at both the far and near ends of the crystal can achieve better three-dimensional position resolution in thick crystals. Studies have shown that using PDPC coupled to a 22-μm-thick LYSO crystal and using the k-value nearest neighbor algorithm, a single-end readout detector can achieve a planar resolution of 1.7 mm and a DOI resolution of 3.7 mm, while dual-end readout increases the resolution to 1.1 mm and 2.4 mm, respectively. It can be seen that in order to simultaneously achieve high three-dimensional positioning accuracy and high detection efficiency of PET detectors, a dual-end readout scheme based on continuous crystals is a very feasible solution.

[0012] Since the scintillation light distribution in the continuous crystal is truncated at the edge of the crystal, the energy resolution, three-dimensional position resolution and other performance at this position deteriorate. Summary of the invention

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

[0014] The technical solution of the present invention is:

[0015] A double-end readout PET detector based on continuous crystal light sharing, characterized in that it includes N continuous scintillation crystals, a photoelectric converter, and a signal processing unit, where N≥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] N pieces of the continuous scintillation crystals are arranged into a plurality of rows to form an integral structure, and the sides of adjacent continuous scintillation crystals are connected by optical glue;

[0018] Both end faces of each of the continuous scintillation crystals are connected to a photoelectric converter through optical silicone 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 signal of the photoelectric converter.

[0020] Furthermore, the side surfaces of the overall structure are coated with a black pigment coating for absorbing scintillation photons reaching the side boundaries of the overall structure.

[0021] Furthermore, the black pigment coating has a thickness of 0.25 mm.

[0022] Furthermore, the photoelectric converter is a silicon photomultiplier tube array, and each of the silicon photomultiplier tube arrays is used to receive scintillation photons reaching one end of a continuous scintillation crystal.

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

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

[0025] Furthermore, the refractive index of the optical adhesive is 1.70.

[0026] Furthermore, the signal processing unit comprises a 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 dual-end readout PET detector, the steps of which include:

[0028] 1) Arrange N blocks of the continuous scintillation crystals in a row along the X direction to form an integral structure; the photoelectric converter is a silicon photomultiplier tube array composed of M×M silicon photomultiplier tube units, each of which is used to receive scintillation photons reaching one end of the continuous scintillation crystal; the nth silicon photomultiplier tube array is used to receive scintillation photons reaching the nth continuous scintillation crystal Block n The 2nth silicon photomultiplier tube array is used to receive the scintillation light from the far end, reaching the nth continuous scintillation crystal block. n The near-end scintillation light, n = 1 ~ N; defines the nth continuous scintillation crystal Block n The middle and far end is the nF end, the nth continuous scintillation crystal Block n The near end is the nf end, and the nth silicon photomultiplier tube array is in Block n The signal received at the nF end in the X direction is recorded as E nFX1 、E nFX2 ,……,E nFXM , the signal received in the Y direction is recorded as E nFY1 、E nFY2 ,……,E nFYM ; The 2nth silicon photomultiplier tube array in Block n The signal received at the nf end in the X direction is recorded as E nfX1 、E nfX2 ,……,E nfXM, the signal received in the Y direction is recorded as E nfY1 、E nfY2 ,……,E nfYM ;

[0029] 2) If the scintillation light is generated by the continuous scintillation crystal blocks at both ends of the overall structure j , j = 1 or N, i.e. the first continuous scintillation crystal Block 1, the Nth continuous scintillation crystal Block N ;Block j The distal end and proximal end of the block are denoted as jF end and jf end respectively. j Adjacent continuous scintillation crystal blocks j’ The distal end and proximal end of Block are denoted as j'F end and j'f end respectively; j Subtract the threshold value from each signal in the X direction of the jF end to get:

[0030]

[0031] And for Block j’ Subtract the threshold value from each signal in the X direction at the j'F end to get:

[0032] β is the set threshold;

[0033] 3) Follow the steps in step 2) to modify the Block j The Y direction signals of the jF end are processed by threshold reduction, and the Block j The jf end performs threshold reduction processing on each signal in the X direction and each signal in the Y direction;

[0034] 4) For the X direction: Use the blocks obtained in steps 2) and 3) j and its adjacent continuous scintillation crystal Block j’ The X-direction signal of the far end is calculated by the centroid method after the threshold is subtracted, and the X-direction far end positioning result is recorded as X jF , using step 2),

[0035] 3) Obtained Block j and its adjacent continuous scintillation crystal Block j’ The near-end X-direction signal after the threshold is subtracted is calculated by the centroid method to obtain the near-end positioning result in the X direction, which is recorded as X jf , and then calculate Block j X-direction positioning result of the inner flashing light j =(X jF +X jf ) / 2; Similarly, we can calculate Blockj The Y direction positioning result of the inner flashing light j , thus getting Block j The location of the internal flashing light (X j , Y j );

[0036] 5) If the scintillation light is generated in the continuous scintillation crystal block in the middle of the overall structure k , the far end and the near end are respectively recorded as kF end and kf end, k = 2 ~ N-1; then according to the Block k Adjacent continuous scintillation crystal blocks k’ 、Block k” Block of each signal in the X direction k The X-direction signal at the kF end is subtracted from the threshold to obtain:

[0037]

[0038] 6) Follow the steps in step 5) to modify the Block k The Y direction signals of the kF end are subjected to threshold reduction processing, and the Block k The signals in the X direction and the Y direction of the kf end are subjected to threshold reduction processing;

[0039] 7) For the X direction: Use the blocks obtained in steps 5) and 6) k and its adjacent continuous scintillation crystal Block k’ and Block k” The far-end X-direction signal after subtracting the threshold is calculated by the centroid method to obtain the far-end positioning result in the X direction, which is recorded 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 near-end X-direction signal after subtracting the threshold is calculated by the centroid method to obtain the near-end positioning result in the X direction, which is recorded as X kf , and then calculate Block k X direction positioning result X k =(X kF +X kf ) / 2; Similarly, we can calculate Block k The Y direction positioning result of the inner flashing light k , thus getting Block k The location of the internal flashing light (X k , Y k ).

[0040] A PET system, characterized by comprising a plurality of the above-mentioned double-end readout PET detectors.

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

[0042] In one embodiment of the present invention, the continuous scintillation crystal is used to block gamma rays, convert the gamma rays into scintillation photons, and transmit them to the photoelectric conversion device.

[0043] Specifically, there is no requirement for the three-dimensional size of the continuous scintillation crystal, and the crystal size is arbitrary. Further, in one embodiment of the present invention, the continuous scintillation crystal uses a continuous LYSO crystal, and the side surfaces are sequentially connected between N pieces of continuous scintillation crystals using a high refractive index optical glue, so that the N pieces of continuous scintillation crystals are spliced ​​into a whole. The present invention does not limit the arrangement of the N pieces of continuous scintillation crystals. According to the user's requirements for the shape and size of the detector, the N pieces of continuous scintillation crystals can be arranged in one row or in multiple rows.

[0044] Furthermore, for the connected N continuous scintillator crystals, all exposed side surfaces (ie, the side surfaces not bonded) are uniformly painted black with black pigment.

[0045] The structure of the continuous scintillation crystals is as follows Figure 2 shown.

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

[0047] Each silicon photomultiplier tube unit corresponds to one energy signal, that is, each silicon photomultiplier tube array corresponds to M 2 Road energy signal.

[0048] The time signal of each silicon photomultiplier tube unit is aggregated into one channel, that is, each silicon photomultiplier tube array corresponds to one channel of time signal.

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

[0050] In one embodiment of the present invention, the signal processing unit is a conventional electronic circuit, including readout electronic circuits, analog electronics and digital electronics.

[0051] The readout electronics circuit is a symmetrical charge distribution circuit, which is a resistor network structure. 2 100 channels of energy signals are converted into 2M channels of output.

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

[0053] The digital electronics is used to integrate the amplified energy waveform to obtain the energy signal E i At the same time, the time signal after timing is sampled as a trigger signal, and finally all energy and time information is transmitted to the computer. The overall structure of the detector is shown in the figure Figure 3 shown.

[0054] A PET system structure composed of the detector of the present invention is as follows Figure 4 shown.

[0055] To achieve the purpose of the present invention, the present invention also provides a calculation method for the fine three-dimensional positioning of a continuous crystal detector. The calculation method takes a detector composed of four continuous scintillation crystals bonded and spliced ​​in a way as an example, and the number of silicon photomultiplier tube arrays used in the detector in this example is M×M. Users can apply this calculation method to detectors with other bonding methods based on this example.

[0056] Step S1: Definition. Figure 5 As shown, define the X direction signal of the A end (far end) in the first continuous scintillation crystal Block1 as E AX1 、E AX2 ,……,E AXM ;Y direction signal is E AY1 、E AY2 ,……,E AYM ; The X direction signal at end a (near end) in Block1 is E aX1 、E aX2 ,……,E aXM ;Y direction signal is E aY1 、E aY2 ,……,E aYMThe same applies to the second continuous scintillation crystal Block 2, the third continuous scintillation crystal Block 3, and the fourth continuous scintillation crystal Block 4.

[0057] Step S2: Channel subtraction threshold.

[0058] For the X direction:

[0059] If the flashing light is generated in Block 1: (Block 4 is similar to Block 1, and Block 4 needs to use the C-end signal of the adjacent Block 3 to process the D-end signal of Block 4);

[0060] After the threshold is subtracted from each signal in the X direction at Block 1 A:

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

[0062] If the flashing light occurs in Block 2: (Block 3 is the same as Block 2);

[0063] After the threshold is subtracted from each signal in the X direction at Block2 B:

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

[0065] For the Y direction:

[0066] Block 1, Block 2, Block 3, and Block 4 have the same logic. Take Block 1 as an example:

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

[0068] Step S3: Calculate the position.

[0069] If the scintillation light is generated by the continuous scintillation crystal blocks at both ends j , j = 1 or N, i.e. the first continuous scintillation crystal Block 1, the Nth continuous scintillation crystal Block N ;

[0070] For the X direction: use Block j and its adjacent continuous scintillation crystal Block j’The far-end X direction signal after subtracting the threshold is used for the centroid method, and the X-direction far-end positioning result is recorded as X jF ;Use Block j and its adjacent continuous scintillation crystal Block j’ The near-end X direction signal after subtracting the threshold is used as the centroid method, and the near-end positioning result in the X direction is recorded as X jf ; Final Block j The X-direction positioning result is X j =(X jF +X jf ) / 2.

[0071] For the Y direction: By analogy with the calculation method for the X direction, we get Block j The Y direction positioning result of the inner flashing light j ; Finally get Block j The location of the internal flashing light (X j , Y j );.

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

[0073] For the X direction: use Block k and its adjacent continuous scintillation crystal Block k’ and Block k” The far-end X direction signal after subtracting the threshold is used for the centroid method, and the X-direction far-end positioning result is recorded as X kF ;Use Block k and its adjacent continuous scintillation crystal Block k’ and Block k” The near-end X direction signal after subtracting the threshold is used for the centroid method, and the near-end positioning result in the X direction is recorded as X kf ; Final Block k The X-direction positioning result is X k =(X kF +X kf ) / 2.

[0074] For the Y direction: use the calculation method of the X direction to calculate the Block k The Y direction positioning result of the inner flashing light k , and finally get Block k The location of the internal flashing light (X k , Y k ).

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

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

[0077] 2) The continuous crystal light sharing scheme used in the present invention enables the scintillation light to pass through the interface between adjacent continuous crystals and be detected by the photodetectors corresponding to the adjacent crystals, thereby realizing partial recovery of the light distribution cutoff at the edge and improving the spatial resolution of the detector at the edge;

[0078] 3) The dual-end readout scheme used in the present invention makes the three-dimensional spatial resolution of the continuous crystal higher, and ensures that a high spatial resolution can still 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 of the present invention can be selected by the user, and the modular design allows free assembly into PET systems with different axial lengths, making it more convenient to use. BRIEF DESCRIPTION OF THE DRAWINGS

[0080] Figure 1 This is a plan view of the detector structure.

[0081] Figure 2 Schematic diagram of the detector crystal splicing structure.

[0082] Figure 3 This is an overall structural diagram of a detector described in the present invention.

[0083] Figure 4 The figure is a structural diagram of a PET system composed of the detector of the present invention.

[0084] Figure 5 A schematic diagram illustrating the method used in calculating the fine three-dimensional spatial positioning of a double-ended readout detector for continuous crystal light sharing. DETAILED DESCRIPTION

[0085] The following description is used to disclose the present invention so that those skilled in the art can implement the present invention. The preferred embodiments described below are only examples, and those skilled in the art can think of other obvious variations. The basic principles of the present invention defined in the following description can be applied to other embodiments, deformation schemes, improved schemes, equivalent schemes, and other technical solutions that do not deviate from the spirit and scope of the present invention.

[0086] Those skilled in the art should understand that, in the disclosure of the present invention, the terms "vertical", "lateral", "up", "down", "front", "back", "left", "right", "vertical", "horizontal", "top", "bottom", "inside", "outside" and the like indicating the orientation or position relationship are based on the orientation or position relationship shown in the drawings, which are only for the convenience of describing the present invention and simplifying the description, and do not indicate or imply that the device or element referred to must have a specific orientation, be constructed and operated in a specific orientation. Therefore, the above terms should not be understood as limiting the present invention.

[0087] It is to 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 an element may be one, while in another embodiment, the number of the element may be multiple, and the term "one" should not be understood as a limitation on the quantity.

[0088] In the description of the present invention, it should be noted that, unless otherwise clearly specified and limited, the terms "installed", "connected", and "connected" should be understood in a broad sense, for example, it can be a fixed connection, a detachable connection, or an integral connection; it can be a mechanical connection, an electrical connection, or mutual communication; it can be a direct connection, or an indirect connection through an intermediate medium, it can be the internal connection of two elements or the interaction relationship between two elements. For ordinary technicians in this field, the specific meanings of the above terms in the present invention can be understood according to specific circumstances.

[0089] In one aspect, the present invention designs a double-end readout PET detector based on continuous crystal light sharing. The detector has high three-dimensional spatial resolution and high detection efficiency. Any number of the detectors are used to form a PET system, which has higher system spatial resolution than the traditional PET system, can eliminate the parallax caused by the DOI effect, and has higher system sensitivity. In addition, the detector adopts a modular design and can be arbitrarily composed of PET systems with different ring diameters and axis lengths, which can be applied to a variety of imaging scenarios and is more convenient to use.

[0090] by Figure 1 Taking the detector corresponding to the typical embodiment shown as an example, the specific implementation scheme of the invention is described:

[0091] The detector is composed of four 40mm*40mm*20mm continuous LYSO crystals spliced ​​together on the side, and the refractive index of the optical glue used for splicing is 1.70; except for the upper and lower bottom surfaces of the spliced ​​crystals, the other exposed surfaces are evenly painted black with black pigment, and the thickness of the black pigment coating is ensured to be about 0.25mm; the upper and lower bottom surfaces of the crystal are coupled to 8 12×12 SiPM arrays through optical silicone grease; the crystal and SiPM array are placed in the aluminum bracket at the front end of the detector, and a layer of rubber buffer material is added between the SiPM array PCB board and the bracket; each SiPM array is connected to the readout electronics board, which is a PCB board with flexible wires, so that the signals read out at both ends are led to the back end.

[0092] It can be understood that in this specific example of the present invention, four continuous LYSO crystals with a size of 40mm*40mm*20mm are used. The user can customize the number of continuous crystals used in the detector according to actual needs. It is only necessary to ensure that the number of continuous crystals is ≥2. This mainly affects the detection area of ​​the detector and the field of view of the system when used in a PET system. The present invention does not impose any restrictions on this.

[0093] Specifically, the crystal thickness D is 1-20 mm, preferably 20 mm, in order to ensure a higher detection efficiency of the detector.

[0094] Specifically, the silicon photomultiplier tube array described in the present invention is composed of a plurality of (1-3)×(1-3) mm 2 The silicon photomultiplier tube array is composed of silicon photomultiplier tubes, and the thickness of the silicon photomultiplier tube array is 1 mm.

[0095] That is to say, the SiPM array is composed of a plurality of unit silicon photomultiplier tubes, and the unit area of ​​the unit silicon photomultiplier tube can be 1×1 mm 2 , can also be 2×2mm 2 , 3×3mm 2 , the present invention is not limited to this.

[0096] In particular, in this typical embodiment of the present invention, the SiPM array may be selected to have a unit size of 3×3 mm. 2 The 12×12 SiPM array has an overall size of 40×40×1mm 3 The SiPM array has the characteristics of small area and small dead zone, which is conducive to ensuring the detection performance of the detector.

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

[0098] Specifically, the present invention describes that "the surfaces of the spliced ​​crystals except the upper and lower bottom surfaces are uniformly painted black with black pigment" to absorb the scintillation photons reaching the side boundaries. The selected black pigment should ensure a high absorption rate of photons in the visible light band. The present invention does not limit the type and model of the selected black pigment.

[0099] Specifically, the upper surface (near end) and the lower surface (distal end) of the connected N continuous scintillation crystals described in the present invention are coupled to the 2N silicon photomultiplier tube arrays through optical silicone grease.

[0100] Specifically, all upper and lower surfaces of the scintillation crystals described in the present invention and the surfaces used for connection with other crystals are polished, and there is no restriction on whether the other side surfaces painted black are polished or not.

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

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

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

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

Claims

1. A double-ended readout PET detector based on continuous crystal light sharing, characterized in that: It includes N continuous scintillation crystals, a photoelectric converter, and a signal processing unit, where N is greater than or equal to 2; The continuous scintillation crystal is used to block gamma rays, convert the gamma rays into scintillation photons and transmit them to the photoelectric converter; N pieces of the continuous scintillation crystals are arranged into a plurality of rows to form an integral structure, and the sides of adjacent continuous scintillation crystals are connected by optical glue; Both end surfaces of each of the continuous scintillation crystals are connected to a photoelectric converter through optical silicone grease, respectively, for generating electrons according to the received scintillation photons and transmitting them to the readout electronics circuit; The signal processing unit is used to process and calculate the signal of the photoelectric converter.

2. The dual-end readout PET detector according to claim 1, characterized in that: The side surfaces of the integral structure are coated with a black pigment coating for absorbing scintillation photons reaching the side boundaries of the integral structure.

3. The dual-end readout PET detector according to claim 2, characterized in that: The thickness of the black pigment coating is 0.25 mm.

4. The dual-end readout PET detector according to claim 1, 2 or 3, characterized in that: The photoelectric converter is a silicon photomultiplier tube array, and each of the silicon photomultiplier tube arrays is used to receive scintillation photons reaching one end of a continuous scintillation crystal.

5. The dual-end readout PET detector according to claim 1, 2 or 3, characterized in that: The photoelectric conversion device is a photomultiplier tube.

6. The dual-end readout PET detector according to claim 1, 2 or 3, characterized in that: The continuous scintillation crystal is a continuous LYSO crystal.

7. The dual-end readout PET detector according to claim 1, 2 or 3, characterized in that: The refractive index of the optical adhesive is 1.

70.

8. The dual-end readout PET detector according to claim 1, 2 or 3, characterized in that: The signal processing unit comprises a readout electronics circuit, analog electronics and digital electronics; or the signal processing unit is an ASIC readout chip.

9. A positioning method based on the dual-end readout PET detector according to claim 1, comprising the steps of: 1) Arrange N blocks of the continuous scintillation crystals in a row along the X direction to form an integral structure; the photoelectric converter is a silicon photomultiplier tube array composed of M×M silicon photomultiplier tube units, each of which is used to receive scintillation photons reaching one end of the continuous scintillation crystal; the nth silicon photomultiplier tube array is used to receive scintillation photons reaching the nth continuous scintillation crystal Block n The 2nth silicon photomultiplier tube array is used to receive the scintillation light from the far end, reaching the nth continuous scintillation crystal block. n The near-end scintillation light, n = 1 ~ N; defines the nth continuous scintillation crystal Block n The middle and far end is the nF end, the nth continuous scintillation crystal Block n The near end is the nf end, and the nth silicon photomultiplier tube array is in Block n The signal received at the nF end in the X direction is recorded as E nFX1 、E nFX2 ,……,E nFXM , the signal received in the Y direction is recorded as E nFY1 、E nFY2 ,……,E nFYM ; The 2nth silicon photomultiplier tube array in Block n The signal received at the nf end in the X direction is recorded as E nfX1 、E nfX2 ,……,E nfXM , the signal received in the Y direction is recorded as E nfY1 、E nfY2 ,……,E nfYM ; 2) If the scintillation light is generated by the continuous scintillation crystal blocks at both ends of the overall structure j , j = 1 or N, i.e. the first continuous scintillation crystal Block 1, the Nth continuous scintillation crystal Block N ;Block j The distal end and proximal end of the block are denoted as jF end and jf end respectively. j Adjacent continuous scintillation crystal blocks j’ The distal end and proximal end of Block are denoted as j'F end and j'f end respectively; j Subtract the threshold value from each signal in the X direction of the jF end to get: And Block j’ Subtract the threshold value from each signal in the X direction at the j'F end to get: β is the set threshold; 3) Follow the steps in step 2) to modify the Block j The Y direction signals of the jF end are subjected to threshold reduction processing; and the j The jf end performs threshold reduction processing on each signal in the X direction and each signal in the Y direction; 4) For the X direction: Use the blocks obtained in steps 2) and 3) j and its adjacent continuous scintillation crystal Block j’ The far-end X-direction signal after the threshold is subtracted is calculated by the centroid method to obtain the far-end positioning result in the X direction, which is recorded as X jF , using the Block obtained in steps 2) and 3) j and its adjacent continuous scintillation crystal Block j’ The near-end X-direction signal after subtracting the threshold is calculated by the centroid method to obtain the near-end positioning result in the X direction, which is recorded as X jf , and then calculate Block j X-direction positioning result of the inner flashing light j =(X jF +X jf ) / 2; Similarly, we can calculate Block j The Y direction positioning result of the inner flashing light j , thus getting Block j The location of the internal flashing light (X j , Y j ); 5) If the scintillation light is generated in the continuous scintillation crystal block in the middle of the overall structure k , the far end and the near end are respectively recorded as kF end and kf end, k = 2 ~ N-1; then according to the Block k Adjacent continuous scintillation crystal blocks k’ 、Block k” Block of each signal in the X direction k The X-direction signal at the kF end is subtracted from the threshold to obtain: 6) Follow the steps in step 5) to modify the Block k The Y direction signals of the kF end are subjected to threshold reduction processing, and the Block k The signals in the X direction and the Y direction of the kf end are subjected to threshold reduction processing; 7) For the X direction: Use the blocks obtained in steps 5) and 6) k and its adjacent continuous scintillation crystal Block k’ and Block k” The far-end X-direction signal after subtracting the threshold is calculated by the centroid method to obtain the far-end positioning result in the X direction, which is recorded 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 near-end X-direction signal after subtracting the threshold is calculated by the centroid method to obtain the near-end positioning result in the X direction, which is recorded as X kf , and then calculate Block k X direction positioning result X k =(X kF +X kf ) / 2; Similarly, we can calculate Block k The Y direction positioning result of the inner flashing light k , thus getting Block k The location of the internal flashing light (X k , Y k ).

10. A PET system, characterized in that: Comprising a plurality of dual-end readout PET detectors as claimed in claim 1.

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