CN109459783B - PET device, multilayer crystal PET detector, electronic readout module and method thereof - Google Patents
PET device, multilayer crystal PET detector, electronic readout module and method thereof Download PDFInfo
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Abstract
The invention provides a PET device, a multilayer crystal PET detector, an electronic readout module of the multilayer crystal PET detector and an electronic readout method, wherein the multilayer crystal PET detector comprises n layers of discrete scintillation crystal arrays and n layers of photoelectric converter arrays, the discrete scintillation crystal arrays and the photoelectric converter arrays are arranged at intervals, each layer of photoelectric converter array comprises m photoelectric converters, the photoelectric converters are used for converting optical signals of visible photons detected by the photoelectric converters to obtain energy signals and time signals, the time signals of the m photoelectric converters in each layer are combined together, the energy signals of the m photoelectric converters in each layer are separately output, but the energy signals of the photoelectric converters in the layers are connected and combined in a one-to-one correspondence manner, so that the multilayer crystal PET detector is provided with m energy channels and n time channels. The invention has the advantages of reduced reading channel, high processing efficiency, simple circuit implementation, DOI capability, cross-layer Compton calibration capability and good time performance.
Description
Technical Field
The present invention relates to the field of positron emission tomography, and in particular, to a Positron Emission Tomography (PET) apparatus, a multilayer crystal PET detector, an electronic readout module of a multilayer crystal PET detector, and an electronic readout method.
Background
Positron Emission Tomography (PET) technology takes metabolic activity in organisms as a detection standard, and has a good effect in early diagnosis of serious diseases. At present, the detector technology which takes discrete crystals as scintillation crystals among the detectors widely used in the field of positron emission imaging is the most mature. The discrete crystal detector uses a scintillation crystal array arranged in two dimensions to couple with a photoelectric converter, a strip-shaped discrete crystal is often coupled on one photoelectric converter, and photon reaction position and Depth (DOI) information is obtained by using methods such as a gravity center method and light sharing.
The PET detector composed of the discrete crystals has the advantages of simple decoding algorithm, light edge effect, high spatial resolution and the like, but because photons received by the traditional discrete crystal detector are reflected in the crystals for multiple times to reach the photoelectric converter, an auxiliary method is needed to obtain reaction depth information of gamma photons.
While a discrete crystal based multilayer crystal PET detector is modeled as shown in figure 1. The system comprises a detector, a photoelectric converter array, a discrete scintillation crystal array, a multilayer crystal PET detector and a single-layer detector, wherein 1 is a detector reading signal flat cable, 2 is a photoelectric converter array, 3 is the discrete scintillation crystal array, 4 is the multilayer crystal PET detector, and 5 is a single-layer detector. The multilayer crystal PET detector shortens the propagation time of gamma photons in the crystal and increases the total length of the crystal, thereby improving the time resolution of the detector and enhancing the interception capability of the gamma photons. However, because of the photoelectric converter array in each layer, the total data volume of the time channel and the energy channel of the multi-layer crystal PET detector is large, which reduces the processing efficiency of the detector to some extent.
Disclosure of Invention
According to one aspect of the present invention, a multi-layered crystal PET detector with high processing efficiency is provided.
The multilayer crystal PET detector comprises n layers of discrete scintillation crystal arrays and n layers of photoelectric converter arrays, wherein the discrete scintillation crystal arrays and the photoelectric converter arrays are arranged at intervals, each layer of photoelectric converter array comprises m photoelectric converters, the photoelectric converters are used for converting optical signals of visible photons detected by the photoelectric converters to obtain energy signals and time signals, the time signals of the m photoelectric converters in each layer are combined together, the energy signals of the m photoelectric converters in each layer are output independently, but the energy signals of the photoelectric converters between the layers are combined in a one-to-one correspondence manner, so that the multilayer crystal PET detector has m energy channels and n time channels.
Preferably, the energy signals of the photoelectric converters between the layers are connected in series in one-to-one correspondence from the first layer to the nth layer.
Preferably, n × m energy signals in the n-layer photoelectric converter array are connected to one hub through a flat cable, and the energy signals of the photoelectric converters between layers on the hub are connected and combined in one-to-one correspondence to obtain m energy channels.
According to another aspect of the present invention, there is provided an electronic readout module for use in the above multilayer crystal PET detector, connected to the photosensor array, the electronic readout module comprising an energy readout circuit for reading energy signals of m energy channels and a time readout circuit for reading time signals of n time channels.
Preferably, the electronic readout module further comprises:
the n time amplifying circuits are connected with the n time channels in a one-to-one correspondence manner;
the n comparators are connected with the n time amplifying circuits in a one-to-one correspondence manner; and
the n time detection circuits are connected with the n comparators in a one-to-one correspondence manner;
the time amplification circuit is used for inputting a time signal to the comparator; the comparator is used for comparing an input time signal with a threshold voltage to obtain an output value and transmitting the output value to the time detection circuit; the time detection circuit is used for inputting an output value to the time readout circuit.
According to another aspect of the invention, there is provided an electronic readout method of the above multilayer crystal PET detector, comprising:
step S100: judging whether cross-layer Compton scattering exists or not, and if not, performing the step S200; if yes, go to step S300;
step S200: judging the number of layers of a detector where the reaction is positioned according to the time signal, and decoding the position and reaction depth information of gamma photons based on the m energy channel information of the layers;
step S300: and comparing the energy of different layers based on the time signal, and then determining the layer where the photon reaction is located according to the result of energy comparison.
Preferably, in step S300, after setting the threshold voltage of the comparator, the output pulse lengths of the comparator may be different based on different time signal responses, and the response position where compton scattering occurs is confirmed by comparing the pulse width or the trigger time of the time signal.
According to another aspect of the present invention, a positron emission tomography apparatus is provided, wherein the positron emission tomography apparatus comprises a data processing module, the above multilayer crystal PET detector, and the above electronic readout module, and the data processing module is connected to the electronic readout module and is configured to perform data processing and image reconstruction on the energy signal and the time signal to obtain a scan image of an object to be imaged.
The invention has the following advantages because the energy channels are combined and the time channels are not combined:
(1) The circuit is simpler, and enough data can be acquired only by providing energy channels with the number of single-layer photoelectric converters and time channels with the number of detector layers.
(2) Without cross-layer compton scattering, the reaction depth can be determined from the time signals of n layers, since there is only one layer of SiPM array time signal per event.
(3) The method has the cross-layer Compton calibration capability, in the case of cross-layer Compton scattering, more than one layer of SiPM arrays have time signals in each event, and the response position of the occurrence of the Compton scattering can be confirmed by comparing the pulse width or the trigger time of the time signals.
(4) By linking the timing signals of all sipms of a single layer together, the timing performance is not degraded compared to a non-hierarchical design.
In conclusion, the invention has the advantages of reduced read channel, high processing efficiency, simple circuit implementation, DOI capability, cross-layer Compton calibration capability and good time performance.
A series of concepts in a simplified form are introduced in the summary of the invention, which is described in further 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.
The advantages and features of the present invention are described in detail below with reference to the accompanying drawings.
Drawings
The following drawings of the invention are included to provide a further understanding of the invention. The drawings illustrate embodiments of the invention and, together with the description, serve to explain the principles of the invention. In the drawings, there is shown in the drawings,
FIG. 1 is a block diagram of a prior art multi-layered crystal PET detector;
FIG. 2 is an exploded view of a multilayer crystal PET detector according to an embodiment of the invention;
FIG. 3 is a schematic diagram of an electronic readout module for use with a multilayer crystal PET detector according to an embodiment of the invention;
FIG. 4 is a schematic diagram of the comparison of different time signal responses with the output pulse length of the comparator;
figure 5 is a schematic diagram of a positron emitting imaging apparatus in accordance with an embodiment of the invention.
Wherein the reference symbols are
1-detector reading signal line
2-photoelectric converter array
3-discrete scintillation crystal array
4-multilayer crystal PET detector
5-Single layer Detector
100-Detector
11-discrete scintillation crystal array
12-photoelectric converter array
200-electronic reading module
21-energy sensing circuit
22 1 22 n-time amplifying circuit
23 1 23 n-comparator
24 1 24 n-time detection circuit
25-time readout circuit
300-data processing module
Detailed Description
In the following description, numerous details are provided to provide a thorough understanding of the present invention. One skilled in the art will recognize, however, that the following description is merely illustrative of a preferred embodiment of the invention and that the invention can be practiced without one or more of these details. In addition, some technical features that are well known in the art are not described in order to avoid confusion with the present invention.
The invention provides a multilayer crystal PET detector, and referring to fig. 2 and 3 in combination, the multilayer crystal PET detector 100 comprises an n-layer discrete scintillation crystal array 11 and an n-layer photoelectric converter array 12, wherein the n-layer discrete scintillation crystal array 11 and the n-layer photoelectric converter array 12 are arranged at intervals up and down in the height direction, namely, as shown in fig. 2, in the height direction, one layer of discrete scintillation crystal array, one layer of photoelectric converter array \8230 \ 8230, and the arrangement of the layers is arranged at intervals.
The multi-layer crystal PET detector shown in fig. 2 includes n layers of detectors, each of the single-layer detectors includes one layer of discrete scintillation crystal array 11 and one layer of photoelectric converter array 12, in terms of the single-layer detector, the discrete scintillation crystal array 11 is formed by coupling a plurality of scintillation crystals (in fig. 2, the number of rows of the discrete scintillation crystal array 11 is a, the number of columns is b, and the number of scintillation crystals is a × b), the coupled scintillation crystal array has an upper surface and a lower surface, the photoelectric converter array 12 is formed by coupling a plurality of photosensors (in fig. 2, the number of rows of the photoelectric converter array 12 is d, the number of columns is c, and the number of photosensors is c × d), the photoelectric converter array 12 is coupled to the upper surface of the discrete scintillation crystal array 11, and each photosensor is coupled with a plurality of scintillation crystals, and is configured to detect visible photons or ultraviolet light generated by the gamma photons reacting with the discrete scintillation crystal array 11, and convert optical signals of the detected visible photons to obtain energy signals and time signals.
Referring again to FIG. 2, the multilayer crystal PET detectors have n, n being greater than 2, with the top detector being the top detector and the bottom detector being the bottom detector. The gamma photons can penetrate through the front n-1 layer to reach the top layer detector and then are intercepted and converted into ultraviolet light or visible light by the scintillation crystal of the top layer detector, can also be directly intercepted and converted into ultraviolet light or visible light by the scintillation crystal of the bottom layer detector, and can also directly penetrate through all discrete scintillation crystal arrays. When Compton scattering is not considered, when gamma photons react in the bottom layer crystal, energy is collected by only a bottom layer detector; when gamma photons react in the m-th layer of crystal, energy is collected by one or only m-th layer of detector modules; by evaluating the detectors that collect energy, the reaction depth of the gamma photons can be determined.
Referring to fig. 3 in combination, in order to improve the processing efficiency of the detector, each layer of the photoelectric converter array includes m (m = c × d) photoelectric converters, the time signals of the m photoelectric converters in each layer are combined together, the energy signals of the m photoelectric converters in each layer are output separately, but the energy signals of the photoelectric converters between the layers are combined in a one-to-one correspondence connection manner, so that the multilayer crystal PET detector has m energy channels and n time channels.
Illustratively, the energy signals of the photoelectric converters between the layers are connected in series from the first layer to the nth layer in a one-to-one correspondence. Taking a 4-layer structure as an example, assuming that there are 3 × 3=9 sensors in each layer, 9 energy signals and 9 time signals are output, and the 9 time signals are directly connected together to form 1 time signal, and the connection method can be as follows:
(a) The signals (9 energy signals and 1 time signal) of the first layer are connected to the second layer through the flat cable, energy signals are combined (two signals are combined and directly connected together according to a one-to-one corresponding relation), and time signals are not combined;
(b) The signals (9 energy signals and 2 time signals) merged at the second layer are connected to the third layer through a flat cable, energy signals are merged (two signals are merged and directly connected together according to a one-to-one correspondence relationship), and time signals are not merged;
(c) The combined signals of the third layer (9 energy signals and 3 time signals) are connected to the fourth layer through a flat cable, the energy signals are combined (two signals are combined and directly connected together according to a one-to-one corresponding relation), and the time signals are not combined;
(d) The combined signals of the fourth layer (9 energy signals and 4 time signals) are sent to a subsequent circuit for energy and time measurement.
Illustratively, n × m energy signals in an n-layer photoelectric converter array are connected to one hub through a flat cable, and the energy signals of the photoelectric converters between the layers on the hub are connected and combined in a one-to-one correspondence to obtain m energy channels. Also taking a 4-layer structure as an example, assuming that each layer has 3 × 3=9 sensors, 9 energy signals and 9 time signals are output, and the 9 time signals are directly connected together to form 1 time signal, and the connection mode can be as follows:
(a) The signals of the 1,2,3,4 layers are connected to a line concentration board through the flat cables. A total of 4 × 9=36 energy signals, and 4 × 1 time signals.
(b) 36 energy signals are combined on the line concentration board (according to a one-to-one correspondence relationship, 4 combinations are directly connected together), and time signals are not combined.
(c) And 9 energy signals and 4 time signals are sent to a subsequent circuit for energy and time measurement.
Referring again to fig. 3, the present invention provides an electronic readout module 200 for use in the above-mentioned multi-layered crystal PET detector 100, and connected to the photosensor array 12, the electronic readout module 200 includes an energy readout circuit 21 and a time readout circuit 25, the energy readout circuit 21 is connected to m energy channels for reading energy signals of the m energy channels, and the time readout circuit 25 is connected to n time channels for reading time signals of the n time channels.
It should be noted that, except that the connection manner between the energy readout circuit 21 and the time readout circuit 25 and the multi-layer crystal PET detector is different from the prior art, other structures are basically the same as the prior art, and thus no further description is given here.
The electronic readout module 200 also illustratively includes n time amplification circuits 22 1 ……、22 n N comparators 23 1 ……、23 n And n time detection circuits 24 1 ……、24 n . n time amplifying circuits 22 1 ……、22 n And n time channels T 1 ……、T n Connecting in a one-to-one correspondence manner; n comparators 23 1 ……、23 n The time amplification circuits are connected with the n time amplification circuits in a one-to-one correspondence manner; n time detection circuits 24 1 ……、24 n And the n comparators are connected in a one-to-one correspondence manner. Each time amplifying circuit is used for inputting a time signal to a comparator connected with the time amplifying circuit; each comparator is used for comparing an input time signal with a threshold voltage thereof to obtain an output value and transmitting the output value to a time detection circuit connected with the comparator; each time detection circuit is for inputting an output value to the time readout circuit 25.
Based on the above-mentioned structure of the electronic readout module 200, the multilayer crystal PET detector includes, when performing electronic readout:
step S100: judging whether cross-layer Compton scattering exists or not, and if not, performing the step S200; if yes, step S300 is carried out, whether cross-layer Compton scattering exists or not can be judged through the energy of the SiPMs in different layers, if only one layer of the SiPM photoelectric detectors has energy signals, the cross-layer Compton scattering phenomenon does not occur, and if two or more layers of the SiPM photoelectric detectors have energy signals, the cross-layer Compton phenomenon exists;
step S200: judging the number of detector layers where the reaction is positioned according to the time signal, and decoding the position and reaction depth information of gamma photons based on the m energy channel information of the layers;
step S300: and comparing the energy of different layers based on the time signal, and then determining the layer where the photon reaction is located according to the result of energy comparison.
Preferably, in step S300, after the threshold voltage of the comparator is set, the output pulse length of the comparator may be different based on different time signal responses, and the response position where compton scattering occurs is determined by comparing the pulse width or the trigger time of the time signal.
Specifically, when the multilayer crystal PET detector performs electronic reading, only one layer of SiPM photoelectric converter sequence can receive a time signal in each event when the multilayer detector array receives gamma photons without cross-layer Compton scattering, so that the number of the detector layers where the reaction is located can be judged according to the time signal, and the position and reaction depth information of the gamma photons can be decoded by using m energy channel information of the layers in combination with a positioning algorithm. When the gamma photon reaction generates cross-layer compton scattering, the reaction depth cannot be directly determined by using an energy signal, but n time channels of photoelectric converters of n-layer detectors are not combined, the time signals can be used for comparing the energy of different layers under the condition of the cross-layer compton scattering, then the layer where the photon reaction is located is determined according to the energy comparison result, referring to fig. 4 in combination, after the threshold voltage of a comparator is set, the output pulse lengths of the comparator can be different due to different time signal responses, and the function of determining the energy of the layers is achieved according to the time signals.
According to another aspect of the present invention, there is provided a positron emission imaging apparatus comprising a data processing module 300, the multi-layered crystal PET detector 100 described above, and the electronic readout module 200 described above. The data processing module 300 is connected to the electronic readout module 200, and is configured to perform data processing and image reconstruction on the energy information and the time information to obtain a scanned image of the object to be imaged. Illustratively, the data processing module 300 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), or a Central Processing Unit (CPU), etc.
The present invention has been illustrated by the above embodiments, but it should be understood that the above embodiments are for illustrative and descriptive purposes only and are not intended to limit the invention to the scope of the described embodiments. Furthermore, 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 may be made in accordance with the teachings of the present invention, which variations and modifications are within the scope of the present invention as claimed. The scope of the invention is defined by the appended claims and equivalents thereof.
Claims (4)
1. An electronic readout module-based electronic readout method, wherein the electronic readout module is used in a multilayer crystal PET detector, the multilayer crystal PET detector includes n layers of discrete scintillation crystal arrays and n layers of photoelectric converter arrays, the discrete scintillation crystal arrays and the photoelectric converter arrays are alternately arranged in a height direction, each layer of the photoelectric converter arrays includes m photoelectric converters, the photoelectric converters are used for converting optical signals of visible photons detected by the photoelectric converters to obtain energy signals and time signals, the time signals of the m photoelectric converters in each layer are combined together, the energy signals of the m photoelectric converters in each layer are separately output, but the energy signals of the photoelectric converters in the layers are combined in a one-to-one correspondence manner, so that the multilayer crystal PET detector has m energy channels and n time channels,
the electronic readout module is connected with the photoelectric converter array, the electronic readout module comprises an energy readout circuit and a time readout circuit, the energy readout circuit is used for reading energy signals of m energy channels, the time readout circuit is used for reading time signals of n time channels,
the electronic readout method includes:
step S100: judging whether cross-layer Compton scattering exists or not, and if not, performing the step S200; if yes, go to step S300;
step S200: judging the number of layers of a detector where the reaction is positioned according to the time signal, and decoding the position and reaction depth information of gamma photons based on the m energy channel information of the layers;
step S300: comparing the energy of different layers based on the time signal, and then determining the layer where the photon reaction is located according to the result of the energy comparison, in step S300, after setting the threshold voltage of the comparator, the output pulse length of the comparator may be different based on different time signal responses, and the response position where compton scattering occurs is determined by comparing the pulse width or the trigger time of the time signal.
2. An electronic readout method according to claim 1, characterized in that energy signals of the photoelectric converters between the layers are connected in series in one-to-one correspondence from the first layer to the n-th layer.
3. The electronic readout method according to claim 1, wherein n x m energy signals in the n-layer photoelectric converter array are connected to one line concentrator via a flat cable, and energy signals of the photoelectric converters between the layers on the line concentrator are connected and combined in one-to-one correspondence to obtain m energy channels.
4. The electronic reading method of claim 1, wherein the electronic reading module further comprises:
the n time amplifying circuits are connected with the n time channels in a one-to-one correspondence manner;
the n comparators are connected with the n time amplifying circuits in a one-to-one correspondence manner; and
the n time detection circuits are connected with the n comparators in a one-to-one correspondence manner;
the time amplification circuit is used for inputting a time signal to the comparator; the comparator is used for comparing an input time signal with a threshold voltage to obtain an output value and transmitting the output value to the time detection circuit; the time detection circuit is used for inputting an output value to the time readout circuit.
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| CN201811155545.4A CN109459783B (en) | 2018-09-30 | 2018-09-30 | PET device, multilayer crystal PET detector, electronic readout module and method thereof |
| PCT/CN2019/102536 WO2020063232A1 (en) | 2018-09-30 | 2019-08-26 | Pet device, multilayer crystal pet detector, and electronic reading module and method therefor |
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| DE102004049677B3 (en) * | 2004-10-12 | 2006-06-14 | Siemens Ag | Detector assembly for use in a combined transmission / emission tomography device |
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| RU2401440C2 (en) * | 2005-08-26 | 2010-10-10 | Конинклейке Филипс Электроникс, Н.В. | High-resolution detector for medical visualisation |
| EP2360493A1 (en) * | 2010-02-15 | 2011-08-24 | Bergen Teknologioverføring AS | Detector arrangement for a tomographic imaging apparatus, particularly for a positron emission tomograph |
| WO2013012809A1 (en) * | 2011-07-15 | 2013-01-24 | Brookhaven Science Associates, Llc | Radiation detector modules based on multi-layer cross strip semiconductor detectors |
| CN103513266B (en) * | 2012-06-21 | 2016-12-28 | 苏州瑞派宁科技有限公司 | Multilayer scintillation crystal and pet detector |
| CN103592671B (en) * | 2012-08-15 | 2016-03-30 | 上海联影医疗科技有限公司 | Scintillation crystal array detector and adopt the PET-MR system of this detector |
| CN104155673B (en) * | 2014-07-21 | 2017-04-12 | 北京永新医疗设备有限公司 | Gamma ray imaging detector and system employing same |
| CN105182396B (en) * | 2015-06-29 | 2018-04-24 | 苏州瑞派宁科技有限公司 | A kind of channel multiplexing method that detector signal is read |
| CN105093258B (en) * | 2015-09-30 | 2018-03-23 | 中派科技（深圳）有限责任公司 | Photon survey front-end circuit |
| CN107320121B (en) * | 2016-04-29 | 2021-06-22 | 上海联影医疗科技股份有限公司 | Positron emission tomography photon detection device |
| US10509135B2 (en) * | 2016-09-09 | 2019-12-17 | Minnesota Imaging And Engineering Llc | Structured detectors and detector systems for radiation imaging |
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