CN116047573A - Detector module for an emission imaging device and emission imaging device - Google Patents

Abstract

The invention provides a detector module for an emission imaging device and the emission imaging device. The detector module includes: the collimating crystal layer comprises a plurality of collimating scintillation crystals, and adjacent collimating scintillation crystals are spaced apart from each other to form a plurality of collimating apertures; the receiving crystal layer comprises a plurality of receiving scintillation crystals, and the collimating crystal layer is positioned in front of the receiving crystal layer; and a photosensor layer optically coupled to the receiving crystal layer, or optically coupled to the collimating crystal layer and the receiving crystal layer. Such a detector module may enable three modality imaging of SPECT, PET, and compton cameras, which may selectively perform any one of the three modality imaging or may simultaneously perform multiple modality imaging. The three modality imaging allows for higher imaging efficiency of the detector module. The detector module can display the information of the target organ more comprehensively and has high sensitivity on the imaging function of the system level.

Description

Detector module for an emission imaging device and emission imaging device

Technical Field

The present invention relates to the field of emission imaging devices, and in particular to a detector module for an emission imaging device and an emission imaging device.

Background

With the increasing level of science and technology, there are increasing means by which people deal with complex conditions. The computed tomography technology is a major breakthrough in the field of nuclear medicine imaging equipment. Emission computed tomography (Emission Computed Tomography, ECT), also known as radionuclide computed tomography, is a imaging technique that can display the distribution and stereoscopic images of radionuclides at various levels in the human body. ECT can detect metabolic and blood flow states of an organ and is a dynamic, functional imaging technique. Currently, common ECT techniques include single photon emission computed tomography (Single Photon Emission Computed Tomography, SPECT), positron emission tomography (Positron Emission Computed Tomography, PET), and compton camera imaging.

However, the existing imaging mode of any mode has own defects, and the information acquired by a single imaging mode has certain limitations and cannot completely reflect the overall characteristics of organisms.

Disclosure of Invention

According to an aspect of the present invention, there is provided a detector module for an emission imaging device, the detector module comprising: the collimating crystal layer comprises a plurality of collimating scintillation crystals, and adjacent collimating scintillation crystals are spaced apart from each other to form a plurality of collimating apertures; the receiving crystal layer comprises a plurality of receiving scintillation crystals, and the collimating crystal layer is positioned in front of the receiving crystal layer; and a photosensor layer optically coupled to the receiving crystal layer, or optically coupled to the collimating crystal layer and the receiving crystal layer, wherein the plurality of receiving scintillation crystals have first projections on planes perpendicular to the front-to-back direction, the plurality of collimating apertures have second projections on the planes, respectively, and the first projections of the plurality of receiving scintillation crystals completely cover the second projections of the plurality of collimating apertures. The detector module with such a structure can realize three modality imaging of SPECT, PET and Compton cameras. The detector module may selectively perform any one of SPECT imaging, PET imaging, and compton camera imaging, or simultaneously perform multiple modalities, depending on the type of tracer injected into the subject. The three modality imaging allows for higher imaging efficiency of the detector module. By identifying and classifying the events of the single gamma photon emitted to the detector module one by one, the detector module can display the information of the target organ more comprehensively and has high sensitivity on the imaging function of the system level. Moreover, the detector module is simple in structure, small in overall occupied space and easy to realize.

The second projection of each collimating aperture is covered by the first projection of the corresponding one or more receiving scintillation crystals, and different collimating apertures correspond to different receiving scintillation crystals. In this way, the collimation accuracy of the detector module is better.

Illustratively, the outer contour of each second projection is aligned with the outer contour of the corresponding one or more first projections; or each second projection is smaller than the corresponding one or more first projections. Therefore, the arrangement of the receiving crystal layer and the collimation crystal layer is more orderly, and the processing is convenient.

Illustratively, adjacent receiving scintillation crystals are spaced apart from one another such that the plurality of collimating scintillation crystals and the plurality of receiving scintillation crystals are staggered. Therefore, the application range of the detector module is wider, and the consumption of scintillation crystal materials can be saved.

The plurality of receiving scintillation crystals are illustratively arranged in a matrix in close proximity to one another such that the plurality of receiving scintillation crystals includes a plurality of first receiving scintillation crystals corresponding to the plurality of collimating scintillation crystals and a plurality of second receiving scintillation crystals corresponding to the plurality of collimating apertures. When a plurality of receiving scintillation crystals are closely arranged in a matrix, the sensitivity of the detector module is higher.

Illustratively, the first plurality of receiving scintillation crystals and the second plurality of receiving scintillation crystals have the same dimensions. Therefore, the arrangement of the scintillation crystals is more neat and consistent with the specification of the scintillation crystals, and the production is convenient.

Illustratively, each of the plurality of first receiving scintillation crystals and the corresponding collimating scintillation crystal are a unitary piece formed from one complete scintillation crystal. The detector module with the integrated piece is simpler in structure and more convenient to manufacture.

The photosensor layer is optically coupled to the back end of the unitary piece and the plurality of second receiving scintillation crystals. Such a detector module may reduce the number of light sensors used and reduce costs.

Illustratively, a face of the unitary piece that is not coupled to the light sensor layer is provided with a light reflective layer facing an interior of the unitary piece; and/or a face of each of the plurality of second receiving scintillation crystals, which is not coupled to the photosensor layer, is provided with a light reflecting layer towards an interior of the second receiving scintillation crystal. The light reflection layer can prevent scintillation light generated by gamma photons captured by the current scintillation crystal from affecting adjacent scintillation crystals. When scintillation light of a single scintillation crystal is detected, the light reflection layer can improve the detection accuracy.

Illustratively, the size of each of the plurality of collimating scintillation crystals is the same as the size of each of the plurality of collimating apertures. Illustratively, a plurality of collimating scintillation crystals alternate with a plurality of collimating apertures along both the row and column directions. Thereby, the collimation accuracy can be improved.

The detector module further includes a light-transmissive plate to which ends of the plurality of collimating scintillation crystals are secured. The light-transmitting plate can be aligned with the direct scintillation crystal to play a role in fixing, so that the collimating crystal layer can be modularized. When the collimating crystal layer and/or the collimating light sensor layer fail, the parts are convenient to be detached and maintained respectively.

Illustratively, the plurality of light-transmitting plates is each inserted into one of the collimating apertures and holds a collimating scintillation crystal that forms the collimating aperture. The light-transmitting plate is arranged, so that the supporting effect of the alignment scintillation crystal is better, the overall structure of the detector module is more stable, and the detector module is convenient to assemble and disassemble.

Illustratively, each collimating aperture has two light-transmitting plates therein, which are respectively secured to both ends of the collimating scintillation crystal forming the collimating aperture. In this way, the size of the collimation apertures can be determined by the light-transmitting plate, and adjacent collimation scintillation crystals can be connected. In addition, the light-transmitting plate is not arranged between the collimating scintillation crystal and the collimating photosensor layer, so that detection of scintillation light in the collimating photosensor layer aligned with the collimating scintillation crystal is prevented from being influenced. And the two light-transmitting plates can seal the collimation holes at two ends, so that dirt is prevented from entering the collimation holes.

Illustratively, the photosensor layer includes a collimating photosensor layer optically coupled to the collimating crystal layer and a receiving photosensor layer optically coupled to the receiving crystal layer. In this way, the collimating crystal layer and the receiving crystal layer are respectively optically coupled to the light sensor layer, so that the detection sensitivity, the collimation accuracy and the spatial resolution of the detector module can be improved.

Illustratively, the collimating light sensor layer is optically coupled to either the front end or the back end of the collimating crystal layer; and/or the receiving light sensor layer is optically coupled to the front or back end of the receiving crystal layer. The relative position relation among the light sensor layer, the collimation crystal layer and the receiving crystal layer can be adjusted according to actual needs, so that the application range of the sensor module is enlarged.

Illustratively, a face of each of the plurality of collimated scintillation crystals that is not coupled to the collimated photosensor layer is provided with a light reflective layer toward an interior of the collimated scintillation crystal. The light reflection layer of the collimation scintillation crystal can prevent scintillation light generated by gamma photons captured by the current scintillation crystal from affecting adjacent scintillation crystals. When scintillation light of a single scintillation crystal is detected, the light reflection layer can improve the detection accuracy.

The face of each of the plurality of receiving scintillation crystals, which is not coupled with the receiving photosensor layer, is provided with a light reflecting layer toward the inside of the receiving scintillation crystal. The light reflection layer of the receiving scintillation crystal can prevent scintillation light generated by gamma photons captured by the current scintillation crystal from affecting adjacent scintillation crystals. When scintillation light of a single scintillation crystal is detected, the light reflection layer can improve the detection accuracy.

Illustratively, the collimating crystal layer is spaced apart from or in close proximity to the receiving crystal layer. The collimating crystal layer is integrally spaced from the receiving crystal layer, thus improving the collimation accuracy of the collimating crystal layer.

According to another aspect of the present invention, there is also provided an emission imaging device. The emission imaging device comprises a detection ring on which a plurality of detector modules of any one of the above are arranged, the collimating crystal layer of the plurality of detector modules is located radially inward of the receiving crystal layer, the plurality of detector modules are arranged in pairs, and each pair of detector modules comprises a first detector module and a second detector module arranged opposite in the radial direction of the detection ring.

Illustratively, in the case that the photosensor layer includes a collimated photosensor layer optically coupled to the collimating crystal layer and a receiving photosensor layer optically coupled to the receiving crystal layer, the transmitting imaging device further includes a readout circuit and a data processing module, the readout circuit being connected to the collimated photosensor layer and the receiving photosensor layer for receiving the electrical signals output by the collimated photosensor layer and the receiving photosensor layer and outputting the energy information and the corresponding time information detected by the collimated photosensor layer and the receiving photosensor layer; the data processing module is connected with the read-out circuit and is used for carrying out data processing and image reconstruction based on the energy information and the time information.

The data processing module is for determining, based on the energy information and the time information, that a single gamma photon incident event detected by the receiving light sensor layer of one of the first detector module and the second detector module that is less than or equal to a first energy threshold and that the energy detected by the collimated light sensor layer of that detector module is zero is a first type of event, and performing SPECT image reconstruction using the first type of event, wherein the first energy threshold is the energy of gamma rays generated by the decay of a tracer employed for SPECT imaging.

The data processing module is for determining, based on the energy information and the time information, that a single gamma photon incident event for which a sum of energies detected by the receiving light sensor layer and the collimated light sensor layer of one of the first detector module and the second detector module is less than or equal to a first energy threshold and the energy detected by the collimated light sensor layer is not zero is a second type event, and performing SPECT image reconstruction using the second type event. The data processing module is configured to determine, based on the energy information and the time information, gamma photons having energies detected by the collimated light sensor layers of the first detector module and the second detector module that are each greater than or equal to a first energy threshold and less than or equal to a second energy threshold, the gamma photons being of a third type of event for incident events, and reconstruct PET images using the third type of event.

The data processing module is used for determining that gamma photons, detected by the light receiving sensor layers of the first detector module and the second detector module, are larger than or equal to the first energy threshold and smaller than or equal to the second energy threshold, are fourth-class events for incident events based on the energy information and the time information, and PET image reconstruction can be carried out by utilizing the fourth-class events.

The data processing module is for determining, based on the energy information and the time information, that gamma photons, each of which is detected by the receiving light sensor layer of one of the first detector module and the second detector module and the collimated light sensor layer of the other of the first detector module and the second detector module as being greater than or equal to the first energy threshold and less than or equal to the second energy threshold, are of a fifth type of events for incident events, and performing PET image reconstruction using the fifth type of events.

The data processing module is for determining, based on the energy information and the time information, that the sum of the energies detected by the collimated light sensor layer and the receiving light sensor layer of the first detector module and the energy detected by the collimated light sensor layer or the receiving light sensor layer of the second detector module are each greater than or equal to the first energy threshold and less than or equal to the second energy threshold are a sixth type of event for the incident event, and PET image reconstruction may be performed using the sixth type of event.

The data processing module is for determining, based on the energy information and the time information, that gamma photons of which the energy detected by the collimated light sensor layer or the receiving light sensor layer of the first detector module and the sum of the energy detected by the collimated light sensor layer and the receiving light sensor layer of the second detector module are both greater than or equal to the first energy threshold and less than or equal to the second energy threshold are a seventh type of event for the incident event, and PET image reconstruction may be performed using the seventh type of event.

The data processing module is for determining, based on the energy information and the time information, that gamma photons of which the sum of energies detected by the collimated light sensor layer and the receiving light sensor layer of the first detector module and the sum of energies detected by the collimated light sensor layer and the receiving light sensor layer of the second detector module are both greater than or equal to a first energy threshold and less than or equal to a second energy threshold are eighth-type events for incident events, and PET image reconstruction may be performed using the eighth-type events.

In the summary, a series of concepts in a simplified form are introduced, which will be further described in detail in the detailed description section. This summary is not intended to identify key features or essential features of the claimed subject matter, nor is it intended to be used as an aid in determining the scope of the claimed subject matter.

Advantages and features of the invention are described in detail below with reference to the accompanying drawings.

Drawings

The following drawings are included to provide an understanding of the invention and are incorporated in and constitute a part of this specification. Embodiments of the present invention and their description are shown in the drawings to explain the principles of the invention. In the drawings of which there are shown,

FIG. 1 is a cross-sectional view of a detector module according to a first exemplary embodiment of the invention;

FIG. 2 is a cross-sectional view of a detector module according to a second exemplary embodiment of the invention;

FIG. 3 is a cross-sectional view of a detector module according to a third exemplary embodiment of the invention;

FIG. 4 is a cross-sectional view of a detector module according to a fourth exemplary embodiment of the invention;

FIG. 5 is a cross-sectional view of a detector module according to a fifth exemplary embodiment of the invention;

FIG. 6 is a cross-sectional view of a detector module according to a sixth exemplary embodiment of the invention;

FIG. 7a is a cross-sectional view of a detector module according to a seventh exemplary embodiment of the invention;

FIG. 7b is a cross-sectional view of a detector module according to an eighth exemplary embodiment of the invention;

FIG. 8a is a top view of a collimating crystal layer of a detector module, according to an exemplary embodiment of the present invention;

FIG. 8b is a cross-sectional view of the collimating crystal layer of FIG. 8 a;

FIG. 9a is a cross-sectional view of a collimating crystal layer of a detector module according to another exemplary embodiment of the present invention;

FIG. 9b is a cross-sectional view of a collimating crystal layer of a detector module according to yet another exemplary embodiment of the present invention;

FIG. 9c is a cross-sectional view of a collimating crystal layer of a detector module, according to yet another exemplary embodiment of the present invention;

FIG. 10a is a cross-sectional view of a detector module according to a ninth exemplary embodiment of the invention;

FIG. 10b is a cross-sectional view of a detector module according to a tenth exemplary embodiment of the invention;

FIG. 10c is a cross-sectional view of a detector module according to an eleventh exemplary embodiment of the invention;

FIG. 11 is a cross-sectional view of a detector module according to a twelfth exemplary embodiment of the invention;

fig. 12 is a schematic diagram of an emissive imaging device according to an exemplary embodiment of the present invention; and

fig. 13 is an enlarged view of a detection ring in the emission imaging device shown in fig. 12.

Detailed Description

In the following description, numerous specific details are set forth in order to provide a more thorough understanding of the present invention. It will be apparent, however, to one skilled in the art that the invention may be practiced without one or more of these details. In other instances, well-known features have not been described in detail in order to avoid obscuring the invention.

In the following description, a detailed structure will be presented for a thorough understanding of the present invention. It will be apparent that embodiments of the invention are not limited to the specific details set forth in the claims. Preferred embodiments of the present invention are described in detail below, however, the present invention may have other embodiments in addition to these detailed descriptions.

The operation of PET, SPECT, and compton imaging devices generally includes: the method comprises the steps of injecting tracers with radioactive substances into a human body, absorbing the tracers by target organs and tissues due to specific molecular structures, enabling unstable radioactive substances in the tracers to decay to emit rays, converting the rays into electric signals after corresponding rays are received by a detector of an imaging device, and finally completing image reconstruction according to a reconstruction algorithm. The three nuclear medicine imaging technologies are dynamic and functional imaging technologies, and can observe information on metabolism, blood flow state and the like of organs.

The PET imaging equipment is mainly used for detecting diseases such as tumors, lungs, nerves, heart organs and the like. The radionuclide labeled with short half-life is used for metabolizing needed nutrients such as glucose, protein and the like, such as 18F-FDG (fluorodeoxyglucose), and the like, and the radionuclide labeled cells are injected into a human body for imaging. Because metabolism is relatively vigorous at the tissue sites of a human body where cancer cells and the like are located, the nutrient substances with radionuclide marks are gathered at the positions of the cancer cells, and positrons generated by decay are combined with negative electrons in surrounding tissues to generate annihilation, so that high-energy gamma photon pairs with energy values of 511keV which are transmitted in opposite directions are generated. The high-energy photons are collected by the scintillation crystal detector and then are converted into visible light signals, the light signals are converted into electric signals through the photoelectric sensor and are sent into the front-end electronic circuit to be processed by the circuit, the gamma photons from the same annihilation event are judged according to the passing time, and the distribution positions of the gamma photons, namely the distribution positions of radiopharmaceuticals, are obtained in a computer through data processing and image reconstruction, so that the metabolic condition and functional information of human organs can be represented.

SPECT imaging equipment is used for detecting and diagnosing diseases of heart, brain, renal artery, thyroid gland, liver and the like. Using a radioisotope as a tracer, such as the usual Tc99m, its decay produces pure gamma rays with a single energy of 141 keV. After the tracer is injected into a human body, the tracer concentrates on the organ to be detected, so that the organ tissue becomes a gamma ray source, and the radiation distribution in the organ tissue is recorded by a detector with a collimator rotating around the human body in vitro. The detector rotates by one angle to obtain a group of data, a plurality of groups of data can be obtained by rotating by one circle, a series of fault plane images with different angles can be established according to the data, and the functional metabolism image of the organism can be obtained by an image reconstruction method to obtain the metabolism condition and functional information of the characterization organ.

Compton imaging devices can be used for radio-tracking imaging, cancer diagnosis, surgical guidance, X-ray fluoroscopic imaging, etc., and are a non-mechanically collimated and non-time-coincident imaging modality that tracks incident photons by means of the physical effects of Compton scattering, i.e., performs "electronic collimation". A typical Compton camera consists of a double-layer detector, incident photons are Compton scattered on the first-layer detector, the detector records the position P1 where reaction occurs and the deposition energy E1, scattered photons are emitted out of the first-layer detector, the scattered photons are completely absorbed at the position R2 on the second-layer detector, the deposition energy is E2, according to the obtained 4 pieces of physical quantity information, the direction of the incident photons can be calculated on conical surfaces with the R1 and R2 as the axis included angle theta, when enough Compton scattering events are detected, a corresponding conical surface can be calculated back for each event, and the conical surface intersection is the spatial position where a radiation source is located in theory, so that imaging is completed.

According to one aspect of the invention, a detector module for an emission imaging device is provided. The detector module includes two scintillation crystal layers. The layer near the entrance face serves as the collimating crystal layer and the layer near the exit face serves as the receiving crystal layer. The collimating crystal layer may act as a collimator when SPECT imaging is performed. Both crystal layers can participate in imaging when PET imaging is performed. The Compton data acquired by the detector module can be used for assisting SPECT imaging and/or PET imaging, so that the sensitivity of the SPECT imaging and/or PET imaging is improved. When the emission imaging equipment adopting the detector module is used for medical imaging, various imaging can be carried out on organ tissues at the same time, for example, the heart valve is imaged by SPECT (single photon emission computed tomography) while the cardiovascular diseases are detected by PET (positron emission tomography) imaging, and the preparation work of the detected object before imaging such as tracer injection, anesthesia and the like is avoided. In addition, when imaging in different modes is performed in multiple times, the gestures of the measured object cannot be identical, and a doctor needs to empirically exclude interference caused by gesture changes and compare states of organ tissues in different images. And the interference of the gesture of the measured object on the imaging result can be avoided when the multi-mode imaging is performed at the same time, so that the diagnosis accuracy is improved. Moreover, different modality imaging can obtain different data of the organ tissue, and the correlation between these data is also very important information when multi-modality imaging is performed simultaneously.

The various detector modules provided by the present invention will be described in detail below with reference to the accompanying drawings. As shown in fig. 1, the detector module 10 may include a collimating crystal layer 100, a photosensor layer 200, and a receiving crystal layer 300.

The collimating crystal layer 100 may include a plurality of collimating scintillation crystals 110. Adjacent collimating scintillation crystals 110 can be spaced apart from one another to form a plurality of collimating apertures 120. A scintillation crystal refers to a crystal that converts the energy of energetic particles into optical energy when a gamma photon incidence event occurs. The collimating scintillation crystal 110 can be lutetium yttrium silicate scintillation crystal (LYSO crystal), bismuth germanate scintillation crystal (BGO crystal), sodium iodide scintillation crystal (NaI crystal), or crystals of a variety of other materials. The collimating scintillation crystal 110 can be prismatic, cylindrical, or a variety of other shapes. The plurality of collimating apertures 120 may have the same size or may have different sizes, that is, the distances between adjacent collimating scintillation crystals 110 may be the same or may be different.

The light sensor layer 200 may be optically coupled to the receiving crystal layer 300, or the light sensor layer 200 may be optically coupled to both the collimating crystal layer 100 and the receiving crystal layer 300. The light sensor layer 200 may include one layer of light sensors or may include multiple layers of light sensors. When the light sensor layer 200 is one layer, it may be optically coupled to the receiving crystal layer 300. When the light sensor layer 200 is a multilayer, it may be optically coupled to the collimating crystal layer 100 and the receiving crystal layer 300, respectively. In this case, the photosensor layer 200 may include a collimated photosensor layer 210 optically coupled to the collimating crystal layer 100 and a receiving photosensor layer 220 optically coupled to the receiving crystal layer 300. The light sensor layer 200 will be described in detail later. One or more photosensors may be included for each photosensor layer 200, which may be of various types, such as photomultiplier tubes (PMTs), silicon photomultiplier tubes (sipms), etc., either existing or as may occur in the future. The scintillation light formed in the scintillation crystal can be detected by the optical sensor coupled with the scintillation crystal, the optical sensor can convert scintillation light signals into electric signals and send the electric signals to the processor at the rear end for data processing and image reconstruction, the distribution position of the radioactive drug is obtained, and metabolic condition and functional information of the organ tissues are further represented.

The collimated light sensor layer 210 can be optically coupled to an end face of the collimating crystal layer 100. The end face may be a light incident end of the collimating crystal layer 100 or a light emitting end of the collimating crystal layer 100. The light incident end refers to an end of the collimating crystal layer 100 where gamma photons are incident, and the light emitting end is opposite to the light incident end. Therefore, the light incident end may also be referred to as a front end, and the light exiting end may also be referred to as a rear end. In the illustrated embodiment, the collimated light sensor layer 210 can be optically coupled to the back end of the collimating crystal layer 100. For photosensors in the collimated photosensor layer 210, one photosensor may be optically coupled to a plurality of the collimated scintillation crystals 110, a plurality of photosensors may be optically coupled to one of the collimated scintillation crystals 110, and one photosensor may be optically coupled to one of the collimated scintillation crystals 110.

The receiving photosensor layer 220 may also be optically coupled to the end face of the receiving crystal layer 300. The end face may be a light incident end of the receiving crystal layer 300 or a light emitting end of the receiving crystal layer 300. Similar to the collimating crystal layer 100, the light incident end of the receiving crystal layer 300 may be a front end thereof, and the light exiting end of the receiving crystal layer 300 may be a rear end thereof. In the illustrated embodiment, the receiving photosensor layer 220 may be optically coupled to the back end of the receiving crystal layer 300. For photosensors in the receiving photosensor layer 220, one photosensor may be optically coupled to a plurality of receiving scintillation crystals 310, a plurality of photosensors may be optically coupled to one receiving scintillation crystal 310, and one photosensor may be optically coupled to one receiving scintillation crystal 310.

It should be noted that the optical sensor layer 200 and the crystal layer optically coupled thereto may be directly coupled by optical glue, or a light guiding layer may be disposed therebetween. The light guiding layer may be constructed of a light guiding material including, but not limited to, a resin light guide, light transmitting glass, liquid light guide, or a variety of other materials.

The receiving crystal layer 300 may include a plurality of receiving scintillation crystals 310. The receiving scintillation crystal 310 can be lutetium yttrium silicate scintillation crystals (LYSO crystals), bismuth germanate scintillation crystals (BGO crystals), sodium iodide scintillation crystals (NaI crystals), or crystals of a variety of other materials. The receiving scintillation crystal 310 can be prismatic, cylindrical, or a variety of other shapes. The receiving scintillation crystal 310 and the collimating scintillation crystal 110 can have the same size, shape, or can have different sizes, shapes, respectively. The adjacent receiving scintillation crystals 310 may be spaced apart from each other or may be closely arranged, and are not particularly limited.

The collimating crystal layer 100 may be positioned in front of the receiving crystal layer 300. The azimuthal terms "front" and "rear" referred to herein and above are both relative to the direction of incidence of the gamma photon, with "front" referring to the side closer to the emission source of the gamma photon; conversely, "rear" refers to the side of the emission source that is remote from the gamma photon. The plurality of receiving scintillation crystals 310 may each have a first projection on a plane perpendicular to the front-to-back direction. That is, each receiving scintillation crystal 310 can have a first projection on that plane. The plurality of collimating apertures 120 may each have a second projection on a plane perpendicular to the front-to-back direction. That is, each of the collimating apertures 120 may have a second projection on the plane. The first projections of the plurality of receiving scintillation crystals 310 can completely cover the second projections of the plurality of collimating apertures 120. On the plane, a first projection of the plurality of receiving scintillation crystals 310 as a whole will completely cover a second projection of the plurality of collimation apertures 120. Each second projection may fall on one or more of the first projections. In particular, for a second projection of a collimation aperture 120, the second projection may be completely covered by a first projection, or the second projection may be completely covered by first projections of multiple receiving scintillation crystals 310; on the other hand, for a first projection that receives the scintillation crystal 310, this first projection may completely cover a second projection of the plurality of collimation apertures 120.

The detector module 10 having such a structure, in the event that a single gamma photon is directed to the detector module 10, if the collimated photosensor layer 210 detects energy deposition, but the receiving photosensor layer 220 does not detect energy deposition, indicates that for this event that gamma photon is directed to the detector module 10, the gamma photon strikes the collimated scintillator crystal 110 and energy deposition occurs; if the collimated photosensor layer 210 detects no energy deposition and the receiving photosensor layer 220 detects energy deposition, it is indicated that for this time the gamma photon is directed to the detector module 10, the gamma photon is not incident on the collimated scintillation crystal 110, but is incident on the receiving scintillation crystal 310 after passing through the collimating aperture 120, and energy deposition occurs, so that this time the gamma photon is considered collimated, and this time the gamma photon is directed to the detector module 10 in accordance with the collimated event; if the collimated photosensor layer 210 detects energy deposition and the receiving photosensor layer 220 also detects energy deposition, it is stated that for this gamma photon event directed to the detector module 10, the gamma photon is incident on the collimated scintillation crystal 110, energy deposition occurs, and Compton scatter occurs, and secondary particles generated by Compton scatter are incident on the receiving scintillation crystal 310, energy deposition also occurs, and this gamma photon event directed to the detector module 10 is considered to be coincident with the Compton scatter event. In the event that a single gamma photon is directed to the detector module 10 as described herein, the gamma photon may be, for example, a single-sided gamma photon generated by the annihilation effect of a positron generated by a metastable technetium-99 (Tc 99 m) decay or by a fluorine-18 (18F) decay. Depending on the detection of energy deposition by the collimated light sensor layer 210 and the receiving light sensor layer 220, it may be identified from a large number of gamma photons directed to the detector module 10 whether a single event corresponds to a collimated event or a Compton scatter event to determine whether the requirements for SPECT imaging or Compton camera imaging, respectively, are met. With the paired detector modules 10, it is possible to determine whether the PET imaging requirements are met from the event that a large number of gamma photons are directed to the detector modules 10 based on the detection of energy deposition by the collimated light sensor layer 210 and the receiving light sensor layer 220.

In particular, the detector module 10 with exemplary embodiments of the present invention achieves a significant level of collimation accuracy in SPECT imaging as compared to prior art SPECT single modality imaging detector modules that utilize lead, tungsten, or the like materials to make the collimator.

As shown in fig. 12, when the detector ring 30 is formed by the detector modules 10, positron generated by decay of PET-imaged tracer causes annihilation with negative electrons in surrounding tissue, producing a pair of gamma photons of 180 ° opposite directions, each of which can be detected by the oppositely disposed detector modules 10. From the energy deposition of the collimated light sensor layer 210 and/or the receiving light sensor layer 220 in the pair of detector modules 10, including time information and energy information, etc., it can be determined whether the event of this gamma photon to the detector module 10 meets the requirements required for PET imaging. In general, scintillation crystals have high efficiency in intercepting gamma photons, and there are few cases where gamma photons pass through the scintillation crystal without energy deposition, and the event that gamma photons pass directly through the scintillation crystal is treated as noise in the detected data.

It follows that a detector module 10 having such a configuration can achieve three modality imaging, SPECT, PET, and compton cameras. The detector module 10 may selectively perform any one of SPECT imaging, PET imaging, and compton camera imaging, or simultaneously perform multiple modalities, depending on the type of tracer injected into the subject. The three modality imaging results in a higher imaging efficiency of the detector module 10. By identifying and classifying the events of a single gamma photon directed to the detector module 10 one by one, the detector module 10 can more fully display information of a target organ and has high sensitivity in a system-level imaging function. Moreover, such a detector module 10 is simple in structure, small in overall occupied space, and easy to implement.

Illustratively, the second projection of each collimating aperture 120 may be covered by the first projection of the corresponding one or more receiving scintillation crystals 310, and different collimating apertures 120 may correspond to different receiving scintillation crystals 310.

Fig. 1-4 illustrate the case where one receiving scintillation crystal 310 corresponds to one collimating aperture 120, that is, the second projection of each collimating aperture 120 is covered by the first projection of the corresponding one receiving scintillation crystal 310. Wherein the receiving scintillation crystal 310 of fig. 1-3 has substantially the same projected area as the corresponding collimating aperture 120, fig. 4 shows that the projected area of the receiving scintillation crystal 310 is larger than the projected area of the corresponding collimating aperture 120. Whereas fig. 2-3 differ from fig. 1 in that the receiving scintillation crystal 310 differs from the collimating scintillation crystal 210 in length. The detector module shown in fig. 1 adopts two scintillation crystals with the same specification to directly combine, and the combination mode has simple structure and less crystal material consumption, but partial gamma rays may not pass through any scintillation crystal or have a short path in the scintillation crystal, so that the probability of being detected is reduced. The case where the gamma ray does not pass through any scintillation crystal is shown by the arrow in fig. 1, for example, a portion of the gamma ray passes only through the gap between the collimation aperture 120 and the adjacent receiving scintillation crystal 310, and the portion of the gamma ray passes through the opposite vertex angle of the two gaps and the incident angle is within the angle alpha indicated by the dashed line in the figure. Fig. 2 and 3 are combinations of short and long crystal arrays, respectively, with fig. 2 increasing the length of the receiving scintillation crystal 310 and fig. 3 increasing the length of the collimating scintillation crystal 110 relative to fig. 1. The angle α becomes smaller in fig. 2 due to the longer length of the receiving scintillation crystal 310, whereby the sensitivity can be improved. In fig. 3, the collimation accuracy is improved because the collimation scintillation crystal 110 is lengthened. But the amount of crystalline material used in the embodiments shown in figures 2 and 3 increases. Fig. 4 illustrates a non-uniform arrangement where the size and spacing of the collimating scintillation crystal 110 and the receiving scintillation crystal 310 are different, which can increase both sensitivity and collimation efficiency, but can present some data processing difficulties. Furthermore, in other embodiments not shown, the collimating scintillation crystal 110 and the receiving scintillation crystal 310 can also be arranged spaced apart along the front-to-back direction. Although the embodiments shown in the figures each show the collimating scintillation crystal 110 and the receiving scintillation crystal 310 closely spaced or with only the photosensor layer in between, the present application does not exclude the case where the collimating scintillation crystal 110 and the receiving scintillation crystal 310 are spaced so much apart that the photosensor layer is not able to occupy the space. Spacing the collimating scintillation crystal 110 from the receiving scintillation crystal 310 can improve the accuracy of collimation, but can lose sensitivity.

Fig. 5 shows the case where one collimating aperture 120 corresponds to a plurality of receiving scintillation crystals 310 (e.g. 2 x 2 receiving scintillation crystals 310), i.e. the second projection of each collimating aperture 120 is covered by the first projection of the corresponding plurality of receiving scintillation crystals 310. By designing the size of the receiving scintillation crystal 310 to be 1/n (n is a positive integer greater than or equal to 2) of the size of the collimation aperture 120, the energy information of gamma rays passing through the collimation aperture 120 in different directions and positions can be resolved to a greater extent, and the resolution of the detector can be improved.

For example, the outer contour of each second projection may be aligned with the outer contour of the corresponding one or more first projections. In the embodiment shown in fig. 1-3, each collimating aperture 120 corresponds to one receiving scintillation crystal 310 in a front-to-back direction (i.e., up-and-down direction as shown in the figures), both having the same cross-sectional area, and the outer contours of both are aligned. Thus, in the aforementioned plane perpendicular to the front-to-back direction, the first projection formed by the receiving scintillation crystal 310 can completely cover the second projection formed by its corresponding collimation aperture 120, and the contour of the first projection is aligned with the contour of the second projection. In the embodiment shown in fig. 5, each collimating aperture 120 corresponds to a plurality (e.g., 2 x 2) of receiving scintillation crystals 310 along the front-to-back direction, the cross-sectional area of the collimating aperture 120 is equal to the cross-sectional area of its corresponding plurality (e.g., 2 x 2) of receiving scintillation crystals 310, and their outer contours are aligned. Thus, in the aforementioned plane perpendicular to the front-to-back direction, the plurality of first projections of the plurality of receiving scintillation crystals 310 can completely cover the second projections of their corresponding collimation apertures 120, and the contours of the plurality of first projections are aligned with the contours of the second projections.

For each of the collimating scintillation crystals 110, it may correspond to one or more receiving scintillation crystals 310, that is, a receiving scintillation crystal 310 is provided behind each of the collimating scintillation crystals 110. For example, fig. 5 shows a case where each of the collimating scintillation crystals 110 corresponds to a plurality of receiving scintillation crystals 310. Fig. 6 shows the case where each of the collimating scintillation crystals 110 corresponds to one of the receiving scintillation crystals 310. In these cases, the receiving scintillation crystals 310 are closely spaced to avoid gamma photons not passing through the scintillation crystals, thus increasing the sensitivity of the detector module 10. But the amount of the scintillator crystal used increases.

The alignment of the outer contour of the second projection of each collimating aperture with the outer contour of the first projection of its corresponding one or more receiving scintillation crystals 310 may result in a more orderly arrangement of the receiving crystal layer 300 and the collimating crystal layer 100, which may facilitate processing, especially if the receiving scintillation crystals 310 within the receiving crystal layer 300 are closely aligned. Moreover, the above-described alignment of the outer contours does not result in overlapping of the receiving scintillation crystal 310 and the collimating scintillation crystal 110, and thus the difficulty in data processing is less.

Optionally, each second projection may be smaller than the corresponding one or more first projections. That is, the cross-sectional area of each of the collimating apertures 120 along the front-to-back direction, corresponding to one or more receiving scintillation crystals 310, is greater than the cross-sectional area of the collimating aperture 120. The second projection of each of the collimating apertures 120 is shown in fig. 4 as being covered by a first projection of one of the receiving scintillation crystals 310. As described above, this may reduce the instances where collimated gamma photons are not captured by the receiving scintillation crystal 310, thus increasing the sensitivity of the detector module 10.

It should be noted that, although fig. 4 shows a case where there is a gap between the receiving scintillation crystals 310, in other embodiments, which are not shown, there may be no gap between the receiving scintillation crystals 310.

2-4, adjacent receiving scintillation crystals 310 can be spaced apart from one another such that the plurality of collimating scintillation crystals 110 and the plurality of receiving scintillation crystals 310 can be staggered. In general, the scintillation crystal has high gamma photon interception efficiency, and the gamma photons pass through the collimating scintillation crystal 110 without energy deposition account for a very small proportion of the whole events, so that such a situation can be considered to be ignored, so as to reduce the usage amount of the scintillation crystal. In the staggered arrangement described herein, the collimating scintillation crystal 110 and the receiving scintillation crystal 310 can have the same or different cross-sectional areas. As shown in fig. 4, the receiving scintillation crystal 310 is slightly thicker than the collimating scintillation crystal 110, so that the sensitivity of the detector module 10 is higher. As shown in fig. 1-3, the receiving scintillation crystal 310 can be completely staggered from the collimating scintillation crystal 110. The plurality of collimating scintillation crystals 110 and the plurality of receiving scintillation crystals 310 are arranged in a staggered manner, and the sizes of the partially collimating scintillation crystals 110 and/or the receiving scintillation crystals 310 can be adjusted to adapt to different requirements, so that the application range of the detector module 10 is wider, and the material consumption of the scintillation crystals can be saved.

For example, as shown in fig. 5-6, the plurality of receiving scintillation crystals 310 can be arranged in a matrix in close proximity to one another such that the plurality of receiving scintillation crystals 310 can include a plurality of first receiving scintillation crystals 311 corresponding to the plurality of collimating scintillation crystals 110 and a plurality of second receiving scintillation crystals 312 corresponding to the plurality of collimating apertures 120. When the plurality of receiving scintillation crystals 310 are closely arranged in a matrix, the sensitivity of the detector module 10 is higher. It should be noted that the second projection of the collimating aperture 120 may have the same size as the first projection of the second receiving scintillation crystal 312 (see fig. 6), or the area of the second projection may be an integer multiple of the area of the first projection (see fig. 5). In other embodiments not shown, the area of the second projection may also be a non-integer multiple of the area of the first projection. For example, the width of the collimating aperture 120 may be 1.5 times the width of the second receiving scintillation crystal 312, and along the width direction (e.g., the horizontal direction in the paper), one collimating scintillation crystal 110 and one collimating aperture 120 may correspond to three receiving scintillation crystals 310, in which case one receiving scintillation crystal 310 may be considered both a first receiving scintillation crystal 311 and a second receiving scintillation crystal 312. The scintillation crystal is typically quadrangular, so that the receiving scintillation crystal 310 and the collimating scintillation crystal 110 can be arranged in a similar manner to the arrangement in the width direction along the direction perpendicular to the paper.

Illustratively, as shown in fig. 6, the plurality of first receiving scintillation crystals 311 and the plurality of second receiving scintillation crystals 312 may have the same dimensions. Like this, the scintillation crystal on the detector module 10 arranges neatly more, and overall structure is simpler, and convenient production easily realizes.

Illustratively, as shown in fig. 7a, each of the plurality of first receiving scintillation crystals 311 and the corresponding collimating scintillation crystal 110 may be a unitary piece 313 formed of one complete scintillation crystal, and the collimating photosensor layer 210 and the receiving photosensor layer 220 may be optically coupled to both ends of the unitary piece 313, respectively. A second receiving scintillation crystal 312 is disposed between two adjacent unitary pieces 313. Although in the illustrated embodiment the number of second receiving scintillation crystals 312 between two adjacent unitary pieces 313 is one, in other embodiments not shown, the number of second receiving scintillation crystals 312 between two adjacent unitary pieces 313 is a plurality. In the event that a single gamma photon is directed to detector module 10, if the gamma photon is incident into unitary piece 313 and energy deposition occurs, such detector module 10 may determine the specific location of the gamma photon incident into unitary piece 313 based on the relative amounts of light signals captured by collimated light sensor layer 210 and received light sensor layer 220 in terms of PET imaging; in SPECT imaging, the event of gamma photon incidence on a body 313 may belong to an uncollimated event, while the event of gamma photon incidence on a second receiving scintillation crystal 312 between two adjacent bodies 313 may correspond to a collimated event. The detector module 10 with the integrated piece 313 has simpler structure and more convenient manufacture.

In the foregoing description, the photosensor layer 200 is a plurality of layers, for example, two layers, namely the collimated photosensor layer 210 and the receiving photosensor layer 220. Alternatively, the photosensor layer 200 may also be one layer, in which case the photosensor layer 200 may be optically coupled to the back ends of the body 313 and the plurality of second receiving scintillation crystals 312 as shown in fig. 7 b. The photosensor layer 200 is optically coupled to all scintillation crystals. In SPECT imaging, during an event in which a single gamma photon is directed to the detector module 10, the detector module 10 determines whether a collimation event is met based on the location and energy information of the light signal captured by the photosensor layer 200. If an event occurs in which a gamma photon is incident on body 313 and energy deposition occurs, the sub-event is considered to be an uncollimated event; if an event occurs in which gamma photons are incident on the receiving scintillation crystal 310 and energy deposition occurs, the sub-event is considered to be coincident with a collimation event. Such a detector module 10 may reduce the number of light sensors used and reduce costs.

For example, the face of the body 313 not coupled with the light sensor layer 200 may be provided with a light reflecting layer towards the inside of the body 313. For example, the face of each of the plurality of second receiving scintillation crystals 312 that is not coupled to the photosensor layer 200 may be provided with a light reflective layer toward the interior of the second receiving scintillation crystal. The light reflection layer can prevent scintillation light generated by gamma photons captured by the current scintillation crystal from affecting adjacent scintillation crystals. When scintillation light of a single scintillation crystal is detected, the light reflection layer can improve the detection accuracy. The light reflecting layer may be formed by spraying, plating (e.g., spraying or plating with a silver film), or pasting a light reflecting material (e.g., ESR reflector). As a high-efficiency reflector, the reflectivity of ESR (Enhanced Specular Reflector) in the whole visible light spectrum range is more than 98%, which is higher than that of other types of reflectors at present. The ESR consists of a high polymer film layer, and is a more environment-friendly reflecting sheet material. The ESR retroreflective sheeting has a thickness of about 40 microns, for example 38 microns.

Illustratively, the size of each of the plurality of collimating scintillation crystals 110 can be the same as the size of each of the plurality of collimating apertures 120. Illustratively, the plurality of collimating scintillation crystals 110 can be arranged alternately with the plurality of collimating apertures 120 along both the row and column directions. As shown in fig. 8a-8b, the size of the collimating scintillation crystal 110 is the same as the size of the collimating aperture 120, and a plurality of collimating scintillation crystals 110 alternate with a plurality of collimating apertures 120 along both the row and column directions. In the collimating crystal layer 100 thus arranged, the collimating apertures 120 may have a uniform size in the row and column directions, and they may be uniformly distributed. Thereby, the collimation accuracy can be improved.

For example, as shown in fig. 9a-9c, the detector module 10 may further comprise a light-transmissive plate 500, and the ends of the plurality of collimating scintillation crystals 110 may be fixed to the light-transmissive plate 500. The light-transmitting plate 500 may act as a fixture for the alignment scintillation crystal 110, enabling the collimating crystal layer 100 to be modularized. In the event of a failure of the collimating crystal layer 100 and/or the collimating light sensor layer 210 optically coupled to the collimating crystal layer 100, it is convenient to disassemble and repair the parts separately. In this case, the alignment photosensor layer 210 may be connected to the light-transmitting plate 500, or may be connected to an end of the alignment scintillator crystal 110 to which the light-transmitting plate 500 is not connected.

Illustratively, the light-transmitting panel 500 may be one or more. When the light-transmitting plate 500 is one, all of the collimating scintillation crystals 110 can be fixed to the light-transmitting plate 500. When the light-transmitting plate 500 is plural, the collimating scintillation crystals 110 may be divided into plural groups, and fixed to the plural light-transmitting plates 500, respectively. Illustratively, as shown in FIG. 9b, where there are a plurality of light-transmitting plates 500, each light-transmitting plate 500 may be inserted into one of the collimating apertures 120 and secured to the collimating scintillation crystal 110 forming that collimating aperture 120. In this case, the size of each light-transmitting plate 500 is the same as the size of the collimating aperture 120, and the edges of the light-transmitting plates 500 may be attached (e.g., bonded) to the sides of the collimating scintillation crystal 110 that form the collimating aperture 120. Thus, the size of the collimation apertures 120 can be determined by the light-transmitting plate 500, and adjacent collimating scintillation crystals 110 can be joined. In addition, the light-transmitting plate 500 is not arranged between the collimating scintillation crystal 110 and the collimating photosensor layer 210, so that detection of scintillation light in the collimating scintillation crystal 110 is prevented from being influenced by the collimating photosensor layer 210.

Within each collimation aperture 120, one or more light-transmitting panels 500 may be provided. In case a light-transmitting plate 500 is provided, it may be located at the exit end, as shown in fig. 9b, at the entrance end, or in the middle of the collimating aperture 120. Where a plurality of light-transmitting plates 500 are disposed within each of the collimating apertures 120, the plurality of light-transmitting plates 500 may be disposed in spaced apart relation to ensure the strength of the connection between the collimating scintillation crystals 110. Illustratively, as shown in fig. 9c, each of the collimating apertures 120 may have two light-transmitting plates 500 therein, and the two light-transmitting plates 500 may be respectively fixed at both ends of the collimating scintillation crystal 110 forming the collimating aperture 120. The light-transmitting plate 500 is arranged in this way, so that the supporting effect of the alignment scintillation crystal 110 is better, the overall structure of the detector module 10 is more stable, and the detector module 10 is convenient to assemble and disassemble. And the two light-transmitting plates 500 can seal the collimating aperture 120 at both ends, preventing contaminants from entering the collimating aperture 120.

10a-10c, light sensor layer 200 may include a collimated light sensor layer 210 optically coupled to collimating crystal layer 100 and a receiving light sensor layer 220 optically coupled to receiving crystal layer 300. In this way, the collimating crystal layer 100 and the receiving crystal layer 300 are optically coupled to the light sensor layer 200, respectively, which may improve the detection sensitivity, collimation accuracy, and spatial resolution of the detector module 10.

As previously described, the collimated light sensor layer 210 can be optically coupled to either the front or back end of the collimating crystal layer 100. Illustratively, the receiving light sensor layer 220 may be optically coupled to either the front end or the back end of the receiving crystal layer 200. The relative positions of collimating crystal layer 100, collimating photosensor layer 210, receiving crystal layer 300, and receiving photosensor layer 220 are shown in three arrangements, for example, in fig. 10a-10 c. In fig. 10a, a collimated light sensor layer 210 is optically coupled to the back end of the collimating crystal layer 100, and a receiving light sensor layer 220 is optically coupled to the back end of the receiving crystal layer 300. In fig. 10b, the collimating photosensor layer 210 is optically coupled to the rear end of the collimating crystal layer 100, and the receiving photosensor layer 220 is optically coupled to the front end of the receiving crystal layer 300, so that the circuit can be simplified, and a circuit board can be used for the collimating photosensor layer 210 and the receiving photosensor layer 220. In fig. 10c, the collimating photosensor layer 210 is optically coupled to the front end of the collimating crystal layer 100, and the receiving photosensor layer 220 is optically coupled to the back end of the receiving crystal layer 300, so that the scintillation crystal space is more compact, and the sensitivity of the detector module 10 can be improved.

Illustratively, the face of each of the collimated scintillation crystals 110 that is not coupled with the collimated photosensor layer 210 may be provided with a light reflective layer toward the interior of the collimated scintillation crystal 110. The light reflecting layer of the collimating scintillation crystal 110 can prevent scintillation light generated by gamma photons captured by the current scintillation crystal from affecting adjacent scintillation crystals. When scintillation light of a single scintillation crystal is detected, the light reflection layer can improve the detection accuracy. The light reflecting layer may be formed by spraying, plating (e.g., spraying or plating with a silver film), or pasting a light reflecting material (e.g., ESR reflector).

For example, the face of each receiving scintillator crystal 310 that is not coupled to the receiving photosensor layer 220 may be provided with a light reflecting layer toward the interior of the receiving scintillator crystal 310. Similar to the light reflecting layer of the collimating scintillation crystal 110, the light reflecting layer of the receiving scintillation crystal 310 can prevent scintillation light generated by gamma photons captured by the current scintillation crystal from affecting neighboring scintillation crystals. When scintillation light of a single scintillation crystal is detected, the light reflection layer can improve the detection accuracy. The light reflecting layer may be formed by spraying, plating (e.g., spraying or plating with a silver film), or pasting a light reflecting material (e.g., ESR reflector).

Illustratively, the collimating crystal layer 100 may be spaced apart from or otherwise in close proximity to the receiving crystal layer 300. As shown in fig. 11, the collimating crystal layer 100 may be spaced apart from the receiving crystal layer 300, thus improving the collimation accuracy of the collimating crystal layer 100. In this case, the collimating crystal layer 100 and the receiving crystal layer 300 have a collimating photosensor layer 210 and a receiving photosensor layer 220, respectively.

According to another aspect of the present invention, as shown in fig. 12, there is provided an emissive imaging device 20, the emissive imaging device 20 may include a detection ring 30. Fig. 13 shows an enlarged view of the probe ring 30. The detector ring 30 may have a plurality of any of the detector modules 10 as described above arranged thereon. Typically, the plurality of detector modules 10 forming one detector ring 30 have substantially the same structure. The present application does not exclude embodiments in which different detector modules 10 form one detector ring 30. A plurality of detection rings 30 may be closely arranged in a direction perpendicular to the paper surface. The detection space enclosed by these detection rings 30 can accommodate the object to be measured. Typically, the detection space may be substantially cylindrical. Of course, embodiments in which the detection space has other shapes are not excluded from the present application. The collimated crystal layer 100 of the plurality of detector modules 10 may be located radially inward of the receiving crystal layer 300. There may be little gap between the light incident ends of the collimating crystal layers 100 of adjacent detector modules 10, thereby enabling all gamma rays generated in the detection space surrounded by the detection ring 30 to be collected by the detector modules 10. The plurality of detector modules 10 may be arranged in pairs, and each pair of detector modules 10 may include a first detector module 301 and a second detector module 302 disposed opposite in a radial direction of the detection ring 30. The pairs of first detector modules 301 and second detector modules 302 may function under PET imaging.

Illustratively, where the photosensor layer 200 includes a collimated photosensor layer 210 optically coupled to the collimating crystal layer 100 and a receiving photosensor layer 220 optically coupled to the receiving crystal layer 300, as shown in fig. 12, the emissive imaging device 20 may further include a readout circuit 40 and a data processing module 50. The readout circuit 40 may be connected to the collimated light sensor layer 210 and the receiving light sensor layer 220, may be used to receive the electrical signals output by the collimated light sensor layer 210 and the receiving light sensor layer 220, and may output the energy information and the corresponding time information detected by the collimated light sensor layer 210 and the receiving light sensor layer 220. Illustratively, the collimated light sensor layer 210 and the receiving light sensor layer 220 of each detector module 10 may be connected to one readout circuit 40. It should be understood that the number of readout circuits 40 and the manner of their connection to the photosensor array 320 can be set as desired, and the invention is not limited thereto. For example, all photosensor arrays 320 of a positron emission imaging device may be connected to the same readout circuitry 40, which readout circuitry 40 may distinguish from which detector module 10 the received electrical signals came to the collimated photosensor layer 210 and/or the receiving photosensor layer 220.

The readout circuitry 40 may read the electrical signals output by the collimated light sensor layer 210 and the receiving light sensor layer 220 and measure the electrical signals to obtain energy information and time information. The energy information may indicate the energy of the visible photons received by the collimated photosensor layer 210 and the receiving photosensor layer 220, respectively. The energy of all visible photons generated by the same gamma photon is accumulated together to be the energy of the gamma photon. The time information may indicate the time of generation of the visible photons received by the photosensor array 320, which may be considered as the arrival time of the gamma photons generating the visible photons to the detector module 10.

The data processing module 50 may be coupled to the readout circuitry 40 for data processing, image reconstruction, etc., based on the energy information and the time information. Thereby, a scanned image of the object to be measured can be obtained. The point of improvement of the present application is not a specific method of data processing and image reconstruction, and will not be described in further detail herein.

Readout circuitry 40 and data processing module 50 may be implemented using any suitable hardware, software, and/or firmware. Illustratively, the data processing module 50 may be implemented using a Field Programmable Gate Array (FPGA), a Digital Signal Processor (DSP), a Complex Programmable Logic Device (CPLD), a Micro Control Unit (MCU), a Central Processing Unit (CPU), or the like.

For ease of description, the principles of three modality imaging are described below by way of example only with respect to the exemplary embodiment shown in fig. 13.

For example, the data processing module 50 may be configured to determine, based on the energy information and the time information, that a single gamma photon incident event detected by the receiving light sensor layer 220 of one of the first detector module 301 and the second detector module 302 that is less than or equal to a first energy threshold and that the energy detected by the collimated light sensor layer 210 of that detector module is zero is a first type event, and may perform SPECT image reconstruction using the first type event, wherein the first energy threshold is an energy of gamma rays generated by decay of a tracer employed for SPECT imaging. Currently, the radionuclide of choice for SPECT imaging is Tc99m, with Tc99m decay producing gamma photon energies of approximately 141keV. Thus, the first energy threshold may be 141keV. The energy information of the first type of event corresponds to the energy carried by the gamma photon emitted when the radionuclide Tc99m selected for SPECT imaging decays, and the energy information is detected on the receiving photosensor layer 220, this event corresponds to a collimation event as described above. The data of the first type of event can be used for SPECT image reconstruction.

In the event that a single gamma photon is directed to the detector module 10, it may also occur that energy is received only in the collimated light sensor layer 210 of the first detector module 301 or the second detector module 302, and that the energy detected by the collimated light sensor layer 210 is less than or equal to the first energy threshold, which is referred to as a zero-th event. The zeroth class of events belong to non-collimated events in SPECT imaging, which may not be considered in SPECT imaging.

Illustratively, the data processing module 50 may be further configured to determine, based on the energy information and the time information, that a single gamma photon incident event for which a sum of energies detected by the receiving light sensor layer 220 and the collimated light sensor layer 210 of one of the first detector module 301 and the second detector module 302 is less than or equal to the first energy threshold and the energy detected by the collimated light sensor layer 210 is not zero is a second type of event, and to reconstruct a SPECT image using the second type of event. The energy information of the second type of event corresponds to the energy carried by gamma photons emitted when the radionuclide Tc99m selected for SPECT imaging decays. The energy information is commonly detected on the collimated light sensor layer 210 and the receiving light sensor layer 220, indicating that a Compton scattering event has occurred. The incidence direction of gamma photons can be deduced from the energy information. When the device needs to have increased sensitivity, a second type of event can be used for SPECT image reconstruction.

For example, the data processing module 50 may be configured to determine, based on the energy information and the time information, that the energy detected by the collimated light sensor layers 210 of the first detector module 301 and the second detector module 302 is greater than or equal to a first energy threshold and less than or equal to a second energy threshold, the gamma photons being of a third type of event for the incident event, and may utilize the third type of event for PET image reconstruction, the first energy threshold being less than the second energy threshold.

Illustratively, the data processing module 50 may be configured to determine, based on the energy information and the time information, that gamma photons, each detected by the received light sensor layers 220 of the first detector module 301 and the second detector module 302, are greater than or equal to the first energy threshold and less than or equal to the second energy threshold, are fourth type of events for the incident event and may utilize the fourth type of events for PET image reconstruction.

For example, the data processing module 50 may be configured to determine, based on the energy information and the time information, that the received light sensor layer 220 of one of the first detector module 301 and the second detector module 302 and the collimated light sensor layer 210 of the other of the first detector module 301 and the second detector module 302 each detect an energy greater than or equal to the first energy threshold and less than or equal to the second energy threshold as a fifth type of event for an incident event, and may utilize the fifth type of event for PET image reconstruction.

The first energy threshold is the energy of gamma rays generated by decay of a tracer adopted by SPECT imaging, and the first energy threshold is the energy of gamma rays generated by decay of the tracer adopted by SPECT imaging. Currently, the radionuclide of choice for SPECT imaging is Tc99m, with Tc99m decay producing gamma photon energies of approximately 141keV. Thus, the first energy threshold may be 141keV. The second energy threshold is the energy of gamma rays generated by decay of the tracer employed for PET imaging. The second energy threshold is the energy of gamma photons generated by decay of the tracer employed for PET imaging. Currently, the radionuclide of choice for PET imaging is 18f, and the energy of the gamma photons produced by 18f decay is approximately 511keV. Thus, the second energy threshold may be 511keV.

The energy information of the third, fourth and fifth events all correspond to the energy carried by gamma photons generated when the radionuclide 18F selected for PET imaging decays, and only one energy signal is detected on the single-sided detector module 10, indicating that compton scattering is not occurring. Thus, the third, fourth, and fifth types of events may be a pair of unscattered gamma photons detected by the detection ring 30, and may be coincident events required for PET imaging. Thus, the third, fourth, and fifth classes of events may be used to perform PET image reconstruction.

Illustratively, the data processing module 50 may be further configured to determine, based on the energy information and the time information, that the sum of the energies detected by the collimated light sensor layer 210 and the receiving light sensor layer 220 of the first detector module 301 and the energy detected by the collimated light sensor layer 210 or the receiving light sensor layer 220 of the second detector module 302 are both greater than or equal to the first energy threshold and less than or equal to the second energy threshold are a sixth type of event for incident events, and may utilize the sixth type of event for PET image reconstruction.

The sixth type of event includes two cases, the first case being that the sum of the energies detected by the collimated light sensor layer 210 and the receiving light sensor layer 220 of the first detector module 301 is greater than or equal to the first energy threshold and less than or equal to the second energy threshold, and the energy detected by the collimated light sensor layer 210 of the second detector module 302 is greater than or equal to the first energy threshold and less than or equal to the second energy threshold; in the second case, the sum of the energies detected by the collimated light sensor layer 210 and the receiving light sensor layer 220 of the first detector module 301 is greater than or equal to the first energy threshold and less than or equal to the second energy threshold, and the energy detected by the receiving light sensor layer 220 of the second detector module 302 is greater than or equal to the first energy threshold and less than or equal to the second energy threshold. The energy information of the sixth type of event corresponds to the energy carried by gamma photons generated when the radionuclide 18F selected for PET imaging decays. The energy information detected by the detection loop 30 indicates that Compton scattering has occurred on the first detector module 301, while unscattered gamma photons are detected on the second detector module 302. For the first detector module 301 where compton scattering occurs, it is necessary to find the initial position of gamma photon incidence through calculation. Thus, it may be a coincidence event required for PET imaging, and thus, a sixth type of event may be used to perform PET image reconstruction.

Illustratively, the data processing module 50 may be further configured to determine, based on the energy information and the time information, that the energy detected by the collimated light sensor layer 210 or the receiving light sensor layer 220 of the first detector module 301 and the sum of the energy detected by the collimated light sensor layer 210 and the receiving light sensor layer 220 of the second detector module 302 are both greater than or equal to the first energy threshold and less than or equal to the second energy threshold are a seventh type of event for incident events, and may utilize the seventh type of event for PET image reconstruction.

The seventh type of event includes two cases, the first case being that the energy detected by the collimated light sensor layer 210 of the first detector module 301 is greater than or equal to the first energy threshold and less than or equal to the second energy threshold, and the sum of the energy detected by the collimated light sensor layer 210 and the receiving light sensor layer 220 of the second detector module 302 is greater than or equal to the first energy threshold and less than or equal to the second energy threshold; in the second case, the energy detected by the receiving light sensor layer 220 of the first detector module 301 is greater than or equal to the first energy threshold and less than or equal to the second energy threshold, and the sum of the energy detected by the collimated light sensor layer 210 and the receiving light sensor layer 220 of the second detector module 302 is greater than or equal to the first energy threshold and less than or equal to the second energy threshold. The energy information of the seventh type of event corresponds to the energy carried by gamma photons generated when the radionuclide 18F selected for PET imaging decays. The energy information detected by the detection loop 30 indicates that Compton scattering has occurred on the second detector module 302, while unscattered gamma photons are detected on the first detector module 301. For the second detector module 302 where compton scattering occurs, the initial position of gamma photon incidence needs to be found through calculation, and the related calculation method is well known to those skilled in the art, and will not be described herein. Thus, it may be a coincidence event required for PET imaging, and thus, a seventh type of event may be used to perform PET image reconstruction.

Illustratively, the data processing module 50 may be further configured to determine, based on the energy information and the time information, that the sum of the energies detected by the collimated light sensor layer 210 and the receiving light sensor layer 220 of the first detector module 301 and the sum of the energies detected by the collimated light sensor layer 210 and the receiving light sensor layer 220 of the second detector module 302 are both greater than or equal to the first energy threshold and less than or equal to the second energy threshold are eighth-type events for incident events, and may utilize the eighth-type events for PET image reconstruction. The energy information of the eighth class of events corresponds to the energy carried by gamma photons generated when the radionuclide 18F selected for PET imaging decays. The energy information detected by the detection loop 30 indicates that Compton scattering has occurred on both the first detector module 301 and the second detector module 302. For the first detector module 301 and the second detector module 302 where compton scattering occurs, it is necessary to find the initial positions of gamma photons incident on the scintillation crystal through calculation and to image the two initial positions by connecting lines. Thus, it may be a coincidence event required for PET imaging, and thus, an eighth class of events may be used for PET image reconstruction.

It will be appreciated that with the emissive imaging device 20 of the present invention, it is possible to determine that the event of a single gamma photon being directed to the detector module 10 corresponds to SPECT imaging, PET imaging, or compton camera imaging based on the energy information and corresponding time information commonly detected by the collimated light sensor layer 210 and/or the receiving light sensor layer 220, wherein more data from compton camera imaging serves to assist in SPECT imaging or PET imaging, improving the sensitivity of SPECT imaging or PET imaging. It follows that such an emissive imaging device 20 can achieve three modality imaging of SPECT, PET, and compton cameras, with high overall device sensitivity, collimation accuracy, and spatial resolution.

The present invention has been illustrated by the above-described embodiments, but it should be understood that the above-described embodiments are for purposes of illustration and description only and are not intended to limit the invention to the embodiments described. In addition, it will be understood by those skilled in the art that the present invention is not limited to the embodiments described above, and that many variations and modifications are possible in light of the teachings of the invention, which variations and modifications are within the scope of the invention as claimed. The scope of the invention is defined by the appended claims and equivalents thereof.

Claims (24)

1. A detector module for an emission imaging device, comprising:

the collimating crystal layer comprises a plurality of collimating scintillation crystals, and adjacent collimating scintillation crystals are spaced apart from each other to form a plurality of collimating apertures;

the receiving crystal layer comprises a plurality of receiving scintillation crystals, and the collimating crystal layer is positioned in front of the receiving crystal layer; and

a photosensor layer optically coupled to the receiving crystal layer, or optically coupled to the collimating crystal layer and the receiving crystal layer, wherein,

the plurality of receiving scintillation crystals respectively have first projections on a plane perpendicular to the front-rear direction, the plurality of collimation holes respectively have second projections on the plane, and the first projections of the plurality of receiving scintillation crystals completely cover the second projections of the plurality of collimation holes.

2. The detector module of claim 1, wherein the second projection of each collimating aperture is covered by the first projection of the corresponding one or more receiving scintillation crystals, and different collimating apertures correspond to different receiving scintillation crystals.

3. The detector module of claim 2, wherein the detector module comprises a plurality of sensors,

The outer contour of each of the second projections is aligned with the outer contour of the corresponding one or more first projections; or alternatively

Each of the second projections is smaller than the corresponding one or more first projections.

4. The detector module of claim 1, wherein adjacent receiving scintillation crystals are spaced apart from one another such that the plurality of collimating scintillation crystals and the plurality of receiving scintillation crystals are staggered.

5. The detector module of claim 1, wherein the plurality of receiving scintillation crystals are arranged in a matrix in close proximity to one another such that the plurality of receiving scintillation crystals includes a plurality of first receiving scintillation crystals corresponding to the plurality of collimating scintillation crystals and a plurality of second receiving scintillation crystals corresponding to the plurality of collimating apertures.

6. The detector module of claim 5, wherein the first plurality of receiving scintillation crystals and the second plurality of receiving scintillation crystals have the same size.

7. The detector module of claim 5, wherein each of the plurality of first receiving scintillation crystals and the corresponding collimating scintillation crystal are a unitary piece formed from one complete scintillation crystal.

8. The detector module of claim 7, wherein the photosensor layer is optically coupled to the unitary piece and a rear end of the plurality of second receiving scintillation crystals.

9. The detector module of claim 8, wherein a face of the unitary piece that is not coupled to the light sensor layer is provided with a light reflective layer toward an interior of the unitary piece; and/or

The face of each of the plurality of second receiving scintillation crystals, which is not coupled to the photosensor layer, is provided with a light reflecting layer towards the interior of the second receiving scintillation crystal.

10. The detector module of claim 1, wherein a size of each of the plurality of collimating scintillation crystals is the same as a size of each of the plurality of collimating apertures.

11. The detector module of claim 10, wherein the plurality of collimating scintillation crystals alternate with the plurality of collimating apertures along both the row and column directions.

12. The detector module of claim 1, further comprising a light-transmissive plate to which ends of the plurality of collimating scintillation crystals are secured.

13. The detector module of claim 12, wherein the plurality of light-transmitting plates each are inserted into one of the collimating apertures and fix the collimating scintillation crystal forming the collimating aperture.

14. The detector module of claim 13, wherein each collimating aperture has two light-transmitting plates secured to each end of the collimating scintillation crystal forming the collimating aperture.

15. The detector module of claim 1, wherein the photosensor layer comprises a collimated photosensor layer optically coupled to the collimating crystal layer and a receiving photosensor layer optically coupled to the receiving crystal layer.

16. The detector module of claim 15, wherein the detector module,

the collimating light sensor layer is optically coupled to the front end or the rear end of the collimating crystal layer; and/or

The receiving light sensor layer is optically coupled to either the front end or the back end of the receiving crystal layer.

17. The detector module of claim 15, wherein the detector module,

a face of each of the plurality of collimated scintillation crystals, which is not coupled with the collimated photosensor layer, is provided with a light reflection layer facing the inside of the collimated scintillation crystal; and/or

The face of each of the plurality of receiving scintillation crystals, which is not coupled with the receiving photosensor layer, is provided with a light reflection layer toward the inside of the receiving scintillation crystal.

18. The detector module of claim 1, wherein the collimating crystal layer is spaced apart or in close proximity to the receiving crystal layer.

19. An emission imaging device comprising a detection ring on which a plurality of detector modules as claimed in any one of claims 1 to 18 are arranged, the collimated crystal layer of the plurality of detector modules being located radially inward of the receiving crystal layer, the plurality of detector modules being arranged in pairs, and each pair of detector modules comprising a first detector module and a second detector module arranged opposite in a radial direction of the detection ring.

20. The emissive imaging device of claim 19, wherein, where the photosensor layer comprises a collimated photosensor layer optically coupled to the collimating crystal layer and a receiving photosensor layer optically coupled to the receiving crystal layer, the emissive imaging device further comprises a readout circuit and a data processing module,

the readout circuit is connected with the light sensor layer and is used for receiving the electric signal output by the light sensor layer and outputting the energy information and the corresponding time information of gamma photons detected by the light sensor layer;

The data processing module is connected with the readout circuit and is used for carrying out data processing and image reconstruction based on the energy information and the time information.

21. The emission imaging device of claim 20, wherein the data processing module is configured to determine, based on the energy information and the time information, that a single gamma photon incident event detected by a receiving light sensor layer of one of the first detector module and the second detector module that is less than or equal to a first energy threshold and that detected by a collimated light sensor layer of the detector module that is zero is a first type of event, and to reconstruct a SPECT image using the first type of event, wherein the first energy threshold is an energy of gamma rays generated by a tracer decay employed in SPECT imaging.

22. The emissive imaging device of claim 20, wherein the data processing module is further configured to determine, based on the energy information and the time information, that a single gamma photon incident event for which a sum of energies detected by a receive light sensor layer and a collimated light sensor layer of one of the first detector module and the second detector module is less than or equal to the first energy threshold and for which the energy detected by the collimated light sensor layer is non-zero is a second type of event, and to reconstruct a SPECT image using the second type of event.

23. The emissive imaging device of claim 20, wherein the data processing module is to:

determining, based on the energy information and the time information, that gamma photons, each of which has an energy detected by a collimated light sensor layer of the first detector module and the second detector module that is greater than or equal to a first energy threshold and less than or equal to a second energy threshold, are a third type of event for incident events, and performing PET image reconstruction using the third type of event, the first energy threshold being less than the second energy threshold; and/or

Based on the energy information and the time information, determining that gamma photons, the energy detected by a light receiving sensor layer of the first detector module and the second detector module is greater than or equal to the first energy threshold and less than or equal to the second energy threshold, are fourth-class events for incident events and performing PET image reconstruction by utilizing the fourth-class events; and/or

Determining, based on the energy information and the time information, that gamma photons, each of which is detected by a receiving light sensor layer of one of the first detector module and the second detector module and a collimated light sensor layer of the other of the first detector module and the second detector module as having an energy greater than or equal to the first energy threshold and less than or equal to the second energy threshold, are a fifth type of event for an incident event, and performing PET image reconstruction using the fifth type of event,

The first energy threshold is energy of gamma rays generated by decay of a tracer adopted by SPECT imaging, and the second energy threshold is energy of gamma rays generated by decay of a tracer adopted by PET imaging.

24. The emissive imaging device of claim 20, wherein the data processing module is further to:

determining, based on the energy information and the time information, that a sum of energies detected by a collimated light sensor layer and a receiving light sensor layer of the first detector module and an energy detected by a collimated light sensor layer or a receiving light sensor layer of the second detector module are both greater than or equal to the first energy threshold and less than or equal to the second energy threshold, gamma photons of incident events being a sixth type of event, and performing PET image reconstruction using the sixth type of event; and/or

Determining, based on the energy information and the time information, that gamma photons of which the sum of the energies detected by the collimated light sensor layer or the receiving light sensor layer of the first detector module and the energies detected by the collimated light sensor layer and the receiving light sensor layer of the second detector module are both greater than or equal to the first energy threshold and less than or equal to the second energy threshold are a seventh type of event for incident events, and performing PET image reconstruction using the seventh type of event; and/or

Based on the energy information and the time information, determining that gamma photons of which the sum of energies detected by the collimation light sensor layer and the receiving light sensor layer of the first detector module and the sum of energies detected by the collimation light sensor layer and the receiving light sensor layer of the second detector module are greater than or equal to the first energy threshold and less than or equal to the second energy threshold are eighth-class events for incident events, and performing PET image reconstruction by utilizing the eighth-class events.

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