A partial-continuous-crystal-based pet detector and pet imaging system

Abstract

The application discloses a PET detector based on partial continuous crystals and a PET imaging system. The PET detector comprises a scintillation crystal array and a single-ended readout photodetector array. The scintillation crystal array is an array formed by arranging multiple composite crystal layers along an X direction, and a first high reflectivity material layer is arranged between adjacent composite crystal layers. Each composite crystal layer comprises a crystal array and a continuous crystal sheet, and the crystal array and the continuous crystal sheet are optically transparently coupled. The crystal array comprises multiple crystal strips arranged along a Y direction, and a second high reflectivity material layer is arranged between adjacent crystal strips. Two end faces of each composite crystal layer along a Z direction are respectively used as an input end and an output end of the scintillation crystal array. The X direction, the Y direction and the Z direction are perpendicular to each other. The output end of the scintillation crystal array is coupled with an input end surface of the photodetector array.

Description

Technical Field

[0001] This invention belongs to the field of gamma-ray imaging detectors, and relates to a PET detector with depth resolution, and more particularly to a PET detector and PET imaging system based on a partially continuous crystal. Background Technology

[0002] Positron emission tomography (PET) utilizes the decay and annihilation of positron-labeled radionuclides within a living organism, producing a pair of back-to-back emitted 511 keV gamma rays. Through coincidence detection and reconstruction algorithms, it can quantitatively image the distribution of the tracer in situ under normal physiological conditions. The core component of PET is the gamma ray detector, typically composed of a scintillation crystal, a light guide, a photodetector, and readout electronics. The scintillation crystal converts the high-energy 511 keV gamma rays into a large number of visible light photons. These photons propagate through the light guide to the photodetector, where they are converted into a large number of electrons through the photoelectric effect and electron multiplication, and finally processed by the readout electronics. The detector's role is to accurately record the location, energy, and timing of gamma ray interaction within the scintillation crystal. For a long time, the Anger centroid method has been used to calculate the two-dimensional interaction location. This method is simple and effective, but it cannot achieve depth-based resolution of the interaction location. If a detector lacks resolution in the depth of interaction (DOI) direction, it will lead to a deterioration in the image resolution of the imaging system, especially for obliquely incident gamma rays. This problem is also known as parallax error. In the field of PET detector research, there are currently a variety of detector designs with DOI resolution, mainly including: phonswich design, offset structure, dual-ended readout, and monolithic crystal. Reference: MOHAMMADI I, CASTRO IFC, CORREIA PMM, et al. Minimization of parallax error in positron emission tomography using depth of interaction capable detectors: methods and apparatus[J]. Biomedical Physics & Engineering Express, 2019, 5(6): 062001.

[0003] Stacked scintillators are one of the most widely used methods, utilizing two or more scintillator materials with different luminescent properties stacked one on top of the other. Waveform discrimination methods can distinguish gamma-effect events occurring in different layers. Its advantages lie in its simplicity and reliability, while its disadvantage is that DOI resolution is limited by the type of scintillator material.

[0004] The offset structure method involves offsetting the crystal arrays of different layers by half a crystal pixel. The top crystal relies on the beam splitting effect of the bottom crystal, and the crystal pixel numbers of different layers can be obtained on a single scatter plot. The advantage of this method is that it does not require additional electronics and the decoding method is simple; the disadvantage is that different DOI layers are obtained on the same scatter plot, which limits the number of layers (usually two).

[0005] Dual-end readout refers to coupling photodetectors to the two end faces of a crystal array, utilizing the diffuse reflection properties between the crystals, and using the signal ratio between the two end faces to obtain high-precision DOI resolution. Currently, dual-end readout offers the best DOI resolution; however, its drawback is that it requires twice the number of photodetectors and electronic channels.

[0006] Continuous crystals refer to photodetectors coupled to a single crystal. They achieve DOI (Domain of Interest) resolution by utilizing the different light distributions of scintillation light emitted from different positions on the photodetector plane. Their advantages include high-precision DOI resolution, eliminating the need for crystal cutting and pixel reflective layers, thus significantly reducing costs and further improving detection efficiency. However, their disadvantages are also obvious: firstly, the increased crystal thickness leads to a significant decrease in 2D position resolution and DOI resolution accuracy; secondly, the resolution deteriorates severely in the crystal edge region (also known as edge effect), and the detector calibration process is complex. These factors make continuous crystal detectors difficult to apply in practice. Researchers at the Institute of Advanced Technology, Chinese Academy of Sciences, proposed a semi-continuous crystal solution, which solved the resolution problem in one of the 2D directions, but the inherent difficulties of continuous crystals still exist.

[0007] Among existing DOI technologies, stacked and offset DOI detectors offer simple decoding methods and are currently the most widely used in PET scanners; however, their DOI resolution remains limited. Detectors based on continuous crystals and dual-end readout can achieve higher DOI resolution, but each has its own significant disadvantages. Therefore, there is currently no complete detector technology solution for high-precision DOI resolution. Summary of the Invention

[0008] In order to overcome the shortcomings of existing continuous crystal detector schemes, such as resolution decrease, edge resolution deterioration and calibration complexity as thickness increases, the present invention aims to provide a PET detector and PET imaging system based on partially continuous crystals.

[0009] The present invention includes a partial continuous crystal detector scheme comprising a scintillation crystal section and a single-end readout silicon photomultiplier tube (SiPM) array. The scintillation crystal section, in the X direction, consists of an array of multiple composite crystal layers separated by a high-reflectivity material. Each composite crystal layer comprises a crystal array and a continuous crystal sheet. The crystal array and the continuous crystal sheet are optically transparent. In the Y direction, the crystal array is formed by bonding N crystal strips together with adhesive and a high-reflectivity material, while the continuous crystal sheet is a single piece of crystal material. A composite crystal layer can be fabricated using two methods: one is to bond the crystal array and the continuous crystal sheet with optically transparent adhesive; in this method, the materials of the crystal array and the continuous crystal sheet need not be the same. The other method is to perform multiple periodic slotting operations on a slightly larger continuous sheet and fill the slots with a high-reflectivity material. Except for the surface coupled to the SiPM array, the entire periphery of the crystal array is encapsulated in a high-reflectivity material. The material coupled to the SiPM array in the scintillation crystal section is optically transparent. The SiPM array consists of multiple SiPMs connected in parallel to form a photodetector plane. One or more crystal strips correspond to one silicon photomultiplier tube.

[0010] The biggest improvement of this invention compared to continuous or semi-continuous crystal schemes lies in the structural design of the scintillation crystal section: it retains the light-blocking structure of traditional array crystals (i.e., the crystal array in the composite crystal layer) while inheriting the characteristics of continuous or semi-continuous crystals (i.e., continuous crystal sheets in the composite crystal layer). This scheme combines the advantages of traditional array schemes and continuous or semi-continuous crystal schemes, solving the difficulties in continuous or semi-continuous crystal applications while maintaining high-precision DOI resolution.

[0011] The detector scheme can obtain the three-dimensional action position of gamma rays within the crystal as follows: Gamma photons interact within the detector crystal and generate scintillation light. First, in the X direction, due to the effect of the highly reflective material between the composite crystal layers, the specific action point can be determined within a certain composite crystal layer i using the Anger centroid method. Then, within this composite crystal layer i, on the one hand, the scintillation light is concentrated by the reflective layer within the crystal array, and the specific crystal strip j in the Y direction can be obtained using the Anger centroid method or its extension. On the other hand, the scintillation light is split by the continuous crystal sheets, forming responses with different distribution broadenings on the photodetector plane. Through a certain light distribution analysis algorithm, the resolution of the DOI direction can be achieved. Optical distribution analysis algorithms can be applied to various methods in continuous or semi-continuous crystal schemes. Some examples are listed below: maximum likelihood function method, reference: Ling, T., et al. (2007). "Depth of interaction decoding of a continuous crystal detector module." Physics in Medicine and Biology 52(8):2213-2228.; neural network algorithm, reference: He, W., et al. (2021). "High-performance coded aperture gamma camera based on monolithic GAGG:Ce crystal." Review of Scientific Instruments 92(1):013106.; K-nearest neighbor method, reference: van Dam, HT, et al. (2011). "A practical method for depth of interaction determination in monolithic scintillator PET detectors." Physics in Medicine and Biology 56(13):4135-4145.; etc.

[0012] This detector design combines the advantages of both crystal arrays and continuous crystals. Even as the crystal thickness increases, the resolution does not decrease rapidly. Edge resolution is improved due to the presence of the crystal array, and detector calibration in both the X and Y directions is simple and reliable.

[0013] The technical solution of this invention is as follows:

[0014] A PET detector based on partially continuous crystals, characterized in that it comprises a scintillation crystal array and a single-ended readout photodetector array; wherein,

[0015] The scintillation crystal array is an array composed of multiple composite crystal layers 1 arranged along the X direction, with a first high-reflectivity material layer 5 between adjacent composite crystal layers; each composite crystal layer includes a crystal array 2 and a continuous crystal sheet 3, and the crystal array 2 and the continuous crystal sheet 3 are optically transparently coupled; the crystal array 2 includes multiple crystal strips arranged along the Y direction, with a second high-reflectivity material layer 4 between adjacent crystal strips; the two end faces of each composite crystal layer along the Z direction serve as the input and output ends of the scintillation crystal array, respectively; the X, Y, and Z directions are mutually perpendicular.

[0016] The output end of the scintillation crystal array is coupled to the input end face of the photodetector array.

[0017] Furthermore, the crystal array 2 and the continuous crystal sheet 3 in the composite crystal layer are bonded together with optically transparent adhesive.

[0018] Furthermore, the composite crystal layer is obtained by repeatedly grooving one side of a continuous crystal sheet and filling the grooves with a high-reflectivity material.

[0019] Furthermore, the continuous crystal sheet 3 is a single piece of crystal.

[0020] Furthermore, the material of the first high reflectivity material layer 5 is barium sulfate, ESR, or titanium dioxide; the composite crystal layer and the first high reflectivity material layer 5 are bonded together with optically transparent adhesive.

[0021] Furthermore, the scintillation crystal array is cuboid; the composite crystal layer has a thickness of 8-30 mm in the Z direction.

[0022] Furthermore, the scintillation crystal array, except for the output end, is encapsulated with a high reflectivity material layer.

[0023] Furthermore, the photodetectors in the photodetector array are silicon photomultiplier tubes, position-sensitive photomultiplier tubes, microchannel plates, avalanche photodiodes, or photomultiplier tube arrays.

[0024] Furthermore, the coupling material between the output end of the scintillation crystal array and the input end face of the photodetector array is an optically transparent material.

[0025] The present invention also provides a PET imaging system, characterized in that it includes the above-described PET detector.

[0026] The advantages of this invention are as follows:

[0027] Compared to stacked scintillators and offset structures, this invention achieves higher DOI resolution, an advantage derived from the continuous crystal wafer. Compared to dual-end readout schemes, this invention uses only single-end readout, significantly reducing the cost of photodetectors and related electronics. Most importantly, compared to continuous crystal schemes, this invention, while maintaining high DOI resolution, addresses the resolution degradation, edge resolution deterioration, and calibration complexity associated with increasing thickness in continuous crystal schemes. It is a highly practical solution capable of simultaneously achieving high detector efficiency and high three-dimensional position resolution. Figure 1 The technical effects of the detector are explained: the composite layer width is 1.5 mm and the groove depth is 1 mm. In the Y direction, the pixel size of the crystal array is 1.5 mm; the two-dimensional position decoding image of the detector obtained by the centroid method shows that the crystal bars are clearly distinguishable. Figure 1 The results show the resolution of the DOI direction of the detector, and the optimal DOI resolution can reach 3.1 mm. Attached Figure Description

[0028] Figure 1 These are renderings of the invention.

[0029] Figure 2 This is a schematic diagram of the structure of a partial continuous crystal PET detector according to the present invention.

[0030] Wherein: 1-composite crystal layer, 2-crystal array in composite layer, 3-continuous crystal sheet in composite layer, 4-reflective layer inside crystal array, 5-high reflectivity material layer between crystal layers, 6-coupling interface between crystal array and continuous crystal sheet, 7-silicon photomultiplier tube array, 8-one pixel of silicon photomultiplier tube. Detailed Implementation

[0031] The present invention will now be described in further detail with reference to the accompanying drawings. The examples given are only for explaining the present invention and are not intended to limit the scope of the present invention.

[0032] Figure 2 This is a preferred embodiment of the present invention. The partial continuous crystal PET detector of the present invention comprises two parts: a scintillation crystal array and a silicon photomultiplier tube array.

[0033] The scintillation crystal array comprises multiple composite crystal layers 1 stacked and coupled in the X direction. The number of these layers is not limited, and they are rectangular in shape, preferably with polished surfaces on all six sides. The thickness of the composite crystal layers in the Z direction is 8-30 mm, and the dimensions in the X and Y directions are not limited and are adjusted according to the dimensions of the silicon photomultiplier tube array.

[0034] The coupling between the composite crystal layers 1 employs a high-reflectivity material layer 5, including but not limited to barium sulfate, enhanced specular reflector (ESR), or titanium dioxide, etc., with a reflectivity generally greater than 90%. This layer isolates scintillation light from different composite crystal layers in the X-direction, enabling better X-direction position resolution using the Anger centroid method. Preferably, the composite crystal layers and the reflective layer 5 are bonded together with optically transparent adhesive; other optically transparent materials can also be used between the composite crystal layers and the high-reflectivity material.

[0035] The composite crystal layer 1 is composed of two parts stacked in the X direction: a crystal array 2 and a continuous crystal sheet 3. It is important to note that the materials of the crystal array 2 and the continuous crystal sheet 3 can be different. The dimensions of the crystal array 2 and the continuous crystal sheet 3 in the X direction are not limited and can be adjusted according to the specific application.

[0036] The composite crystal layer 1 can be fabricated using two methods: one is to bond the crystal array 2 to the continuous crystal sheet 3 with optically transparent adhesive; the other is to make continuous grooves from a slightly larger continuous crystal sheet and fill the grooves with a high-reflectivity material. Other methods capable of forming a similar structure to the composite crystal layer 1 should also be included within the scope of this invention.

[0037] The crystal array 2 can be considered as being composed of multiple coupled crystal strips. The number of crystal strips and the polishing method are unlimited. The internal reflective layer 4 of the crystal array is a high reflectivity material layer. The function of the internal reflective layer 4 is twofold: firstly, to partially isolate the scintillation light in adjacent crystal strips in the crystal array 2 in the Y direction; and secondly, to partially separate the scintillation light in continuous crystal sheets 3 in the Y direction, so as to obtain better position resolution in the Y direction using the energy weighting method.

[0038] The continuous crystal sheet 3 is an uncut, whole scintillation crystal, and the polishing method of its surface is not limited.

[0039] The entire outer surface of the scintillation crystal array, except for the surface coupled to the silicon photomultiplier tube array, is encapsulated with a high-reflectivity material.

[0040] The silicon photomultiplier array 7 is composed of multiple single-pixel silicon photomultiplier tubes arranged in parallel, and the number of single pixels is not limited; the photodetector does not have to be a silicon photomultiplier tube, but can be any photodetector device with position sensitivity, including but not limited to: position-sensitive photomultiplier tubes, microchannel plates, avalanche photodiodes, silicon photomultiplier tube arrays and photomultiplier tube arrays.

[0041] The position of the silicon photomultiplier tube 7 is the same whether it is on the bottom or top surface of the crystal array. The normal to the top surface is along... Figure 2 In the Z-direction.

[0042] The scintillation crystal array and the silicon photomultiplier tube 7 are coupled using optically transparent materials, including but not limited to: optical silicone grease, optical glass, optical silicone oil, etc.

[0043] In summary, the PET detector provided in this embodiment combines the advantages of traditional crystal arrays and continuous crystal schemes. The two-dimensional positioning of the gamma ray interaction point with the crystal in the XY direction can be obtained using the centroid method, while the positioning in the depth of action (DOI) direction is achieved through the beam splitting effect of the continuous crystal sheet. Compared to continuous or semi-continuous crystal schemes, the partially continuous crystal scheme greatly simplifies detector calibration, the resolution does not rapidly deteriorate with increasing crystal thickness, and edge effects are also resolved. Compared to dual-end readout schemes, the partially continuous crystal scheme uses single-end readout, significantly reducing detector costs. Compared to stacked or offset structure schemes, this scheme offers higher DOI resolution.

[0044] Although specific embodiments of the invention have been disclosed for illustrative purposes to aid in understanding and implementing the invention, those skilled in the art will understand that various substitutions, variations, and modifications are possible without departing from the spirit and scope of the invention and the appended claims. Therefore, the invention should not be limited to the content disclosed in the preferred embodiments, and the scope of protection claimed by the invention is defined by the claims.

Claims

1. A PET detector based on partially continuous crystal, characterized in that, This includes scintillation crystal arrays and single-ended readout photodetector arrays; among which, The scintillation crystal array is an array composed of multiple composite crystal layers (1) arranged along the X direction, with a first high reflectivity material layer (5) between adjacent composite crystal layers; each composite crystal layer includes a crystal array (2) and a continuous crystal sheet (3), the crystal array (2) and the continuous crystal sheet (3) are optically transparently coupled, and the crystal array (2) and the continuous crystal sheet (3) are arranged back and forth along the X direction; the crystal array (2) includes multiple crystal strips arranged along the Y direction, with a second high reflectivity material layer (4) between adjacent crystal strips; the two end faces of each composite crystal layer along the Z direction serve as the input and output ends of the scintillation crystal array, respectively; the X, Y, and Z directions are mutually perpendicular; The output end of the scintillation crystal array is coupled to the input end face of the photodetector array; wherein, the crystal array (2) is used to partially isolate the scintillation light through the second high reflectivity material layer (4) to provide position resolution in the Y direction; the continuous crystal sheet (3) is used to split the scintillation light to provide depth resolution in the Z direction.

2. The PET detector according to claim 1, characterized in that, The crystal array (2) in the composite crystal layer is bonded to the continuous crystal sheet (3) with optically transparent adhesive.

3. The PET detector according to claim 1, characterized in that, The composite crystal layer is obtained by repeatedly grooving one side of a continuous crystal sheet and filling the grooves with a high-reflectivity material.

4. The PET detector according to claim 1, 2, or 3, characterized in that, The continuous crystal sheet (3) is a single piece of crystal.

5. The PET detector according to claim 1, 2, or 3, characterized in that, The first high reflectivity material layer (5) is made of barium sulfate, ESR or titanium dioxide; the composite crystal layer and the first high reflectivity material layer (5) are bonded together with optically transparent adhesive.

6. The PET detector according to claim 1, 2, or 3, characterized in that, The scintillation crystal array is cuboid; the composite crystal layer has a thickness of 8-30 mm in the Z direction.

7. The PET detector according to claim 1, 2, or 3, characterized in that, The scintillation crystal array, except for the output end, is encapsulated with a high reflectivity material layer.

8. The PET detector according to claim 1, 2, or 3, characterized in that, The photodetectors in the photodetector array are silicon photomultiplier tubes, position-sensitive photomultiplier tubes, microchannel plates, avalanche photodiodes, or photomultiplier tube arrays.

9. The PET detector according to claim 1, 2, or 3, characterized in that, The coupling material between the output end of the scintillation crystal array and the input end face of the photodetector array is an optically transparent material.

10. A PET imaging system, characterized in that, Includes the PET detector described in any one of claims 1 to 8.

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