WO2016074560A1 - 光电转换器、探测器及扫描设备 - Google Patents

光电转换器、探测器及扫描设备 Download PDF

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WO2016074560A1
WO2016074560A1 PCT/CN2015/092833 CN2015092833W WO2016074560A1 WO 2016074560 A1 WO2016074560 A1 WO 2016074560A1 CN 2015092833 W CN2015092833 W CN 2015092833W WO 2016074560 A1 WO2016074560 A1 WO 2016074560A1
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detector
photoelectric converter
silicon photomultiplier
pet
light guide
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PCT/CN2015/092833
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English (en)
French (fr)
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谢庆国
朱俊
牛明
王璐瑶
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苏州瑞派宁科技有限公司
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Priority to EP15859308.7A priority Critical patent/EP3220425B1/en
Priority to JP2017525003A priority patent/JP7043258B2/ja
Priority to US15/525,266 priority patent/US11022703B2/en
Publication of WO2016074560A1 publication Critical patent/WO2016074560A1/zh

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    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
    • G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
    • G01T1/16—Measuring radiation intensity
    • G01T1/161—Applications in the field of nuclear medicine, e.g. in vivo counting
    • G01T1/164—Scintigraphy
    • G01T1/1641—Static instruments for imaging the distribution of radioactivity in one or two dimensions using one or several scintillating elements; Radio-isotope cameras
    • G01T1/1642—Static instruments for imaging the distribution of radioactivity in one or two dimensions using one or several scintillating elements; Radio-isotope cameras using a scintillation crystal and position sensing photodetector arrays, e.g. ANGER cameras
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
    • G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
    • G01T1/16—Measuring radiation intensity
    • G01T1/20—Measuring radiation intensity with scintillation detectors
    • G01T1/2002—Optical details, e.g. reflecting or diffusing layers
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
    • G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
    • G01T1/16—Measuring radiation intensity
    • G01T1/24—Measuring radiation intensity with semiconductor detectors
    • G01T1/248—Silicon photomultipliers [SiPM], e.g. an avalanche photodiode [APD] array on a common Si substrate
    • H—ELECTRICITY
    • H01—ELECTRIC ELEMENTS
    • H01L—SEMICONDUCTOR DEVICES NOT COVERED BY CLASS H10
    • H01L31/00—Semiconductor devices sensitive to infrared radiation, light, electromagnetic radiation of shorter wavelength or corpuscular radiation and specially adapted either for the conversion of the energy of such radiation into electrical energy or for the control of electrical energy by such radiation; Processes or apparatus specially adapted for the manufacture or treatment thereof or of parts thereof; Details thereof
    • H01L31/02—Details
    • H01L31/0232—Optical elements or arrangements associated with the device
    • H01L31/02327—Optical elements or arrangements associated with the device the optical elements being integrated or being directly associated to the device, e.g. back reflectors
    • H—ELECTRICITY
    • H01—ELECTRIC ELEMENTS
    • H01L—SEMICONDUCTOR DEVICES NOT COVERED BY CLASS H10
    • H01L31/00—Semiconductor devices sensitive to infrared radiation, light, electromagnetic radiation of shorter wavelength or corpuscular radiation and specially adapted either for the conversion of the energy of such radiation into electrical energy or for the control of electrical energy by such radiation; Processes or apparatus specially adapted for the manufacture or treatment thereof or of parts thereof; Details thereof
    • H01L31/08—Semiconductor devices sensitive to infrared radiation, light, electromagnetic radiation of shorter wavelength or corpuscular radiation and specially adapted either for the conversion of the energy of such radiation into electrical energy or for the control of electrical energy by such radiation; Processes or apparatus specially adapted for the manufacture or treatment thereof or of parts thereof; Details thereof in which radiation controls flow of current through the device, e.g. photoresistors
    • H01L31/10—Semiconductor devices sensitive to infrared radiation, light, electromagnetic radiation of shorter wavelength or corpuscular radiation and specially adapted either for the conversion of the energy of such radiation into electrical energy or for the control of electrical energy by such radiation; Processes or apparatus specially adapted for the manufacture or treatment thereof or of parts thereof; Details thereof in which radiation controls flow of current through the device, e.g. photoresistors characterised by potential barriers, e.g. phototransistors
    • H01L31/101—Devices sensitive to infrared, visible or ultraviolet radiation
    • H01L31/102—Devices sensitive to infrared, visible or ultraviolet radiation characterised by only one potential barrier
    • H01L31/107—Devices sensitive to infrared, visible or ultraviolet radiation characterised by only one potential barrier the potential barrier working in avalanche mode, e.g. avalanche photodiodes

Definitions

  • the present invention relates to the field of nuclear medicine imaging technology, and in particular to a photoelectric converter, a detector having the same, and a scanning device having the same.
  • Positron Emission Tomography (PET) detectors are key devices in PET imaging equipment. Their main function is to obtain position, time and energy information during gamma photon deposition in PET systems. The performance of the PET detector directly determines the performance of the entire PET imaging system. In order to improve the imaging performance of the system, it is hoped that the positron emission tomography detector used has high spatial resolution, good time resolution, good energy resolution, High count rate and other characteristics.
  • the function of the photoelectric converter in the PET detector is to convert the scintillation photons outputted by the front-end scintillation crystal into corresponding electric pulses, and multiply and amplify to obtain an electric pulse signal that can be processed by the electronic system.
  • Existing PET detectors, in which photoelectric conversion devices generally use photomultiplier tubes, avalanche photodiodes, position sensitive photomultiplier tubes.
  • photomultiplier tube photocathode by the scintillation photons into photoelectrons, and the photoelectrons multiplied by the pole of a plurality of multiplication, ultor output pulse signal from the photomultiplier tube gain is usually about 106.
  • the photomultiplier tube has the advantages of high gain, low noise, fast time response, etc., which has led to the selection of photomultiplier tubes for most clinical PET photoelectric conversion devices.
  • its volume is generally large, which may limit the spatial resolution of PET detectors and the flexibility of structural design of PET systems; photomultiplier tubes cannot work properly in magnetic fields, and are difficult to be used as photoelectric conversion devices for PET/MRI dual-mode imaging systems.
  • the avalanche photodiode first converts the scintillation photon into photoelectron through the photocathode, and uses the avalanche effect of the photodiode to multiply the photoelectron to obtain an electric pulse signal.
  • the avalanche photodiode can work normally in the magnetic field, and the system performance in PET/MRI dual-mode imaging A certain potential.
  • Avalanche photodiodes are small in size and can be used to design PET detectors with better spatial resolution. Some have adopted avalanche photodiodes to design small animal PETs with higher spatial resolution requirements (Mé lanie Bergeron, Jules Cadorette, Jean- Beaudoin, Martin D.
  • the position-sensitive photomultiplier tube has the advantages of all photomultiplier tubes and can achieve high spatial resolution.
  • the research group has realized the small animal PET with higher spatial resolution by position-sensitive photomultiplier tube (Qingguo).
  • Qingguo position-sensitive photomultiplier tube
  • position-sensitive photomultiplier tubes are very expensive and increase the cost of PET systems.
  • the silicon photomultiplier tube consists of an avalanche photodiode micro pixel cell array operating in Geiger mode with a gain of 106 , and the gain and photomultiplier tube. Comparable, low noise, small size, insensitive to magnetic fields, and good time performance, low cost in mass production, suitable for building PET detectors (Qingguo Xie, Robert G. Wagner, Gary Drake, Patrick DeLurgio, Yun Dong Chin-Tu Chen, Chien-Min Kao, "Performance Evaluation of Multi-Pixel Photon Counters for PET Imaging," in Conference Record of the 2007 IEEE Nuclear Science Symposium, vol. 2, pp. 969-974, 2007).
  • the silicon photomultiplier tube is used as the photoelectric conversion device. Compared with the traditional bulk photomultiplier tube, the silicon photomultiplier tube is compact and compact, and is suitable for constructing a high spatial resolution PET detector, and finally improves the spatial resolution of the entire PET system; The small size of the silicon photomultiplier tube makes it easy to build a variety of detector structures, and is very suitable for building PET detectors with Depth of Interaction (DOI) detection capability, which is greatly improved compared with the larger photomultiplier tube.
  • DOE Depth of Interaction
  • the flexibility of PET system structure construction silicon photomultiplier tube has good time performance, can build Time-of-Flight (TOF)-capable PET detector to improve PET image quality; silicon photomultiplier tube gain High, the working state is not affected by the magnetic field, it is the best choice for the core photoelectric conversion device in the PET/MRI scheme.
  • the cost of the photomultiplier tube, especially the position-sensitive photomultiplier tube, has been high, and the silicon photomultiplier tube is inexpensive in mass production, which can greatly reduce the cost of the PET system.
  • the silicon photomultiplier tube has a gain comparable to that of the photomultiplier tube, and has the advantages of small volume of the avalanche photodiode and insensitivity to the magnetic field, and its time performance is good.
  • the silicon photomultiplier tube has the advantages of a photomultiplier tube and an avalanche photodiode. If the positron emission tomography detector uses a silicon photomultiplier tube as a photoelectric conversion device and can achieve position resolution better, it can be based on a silicon photomultiplier tube. Design and build a low-cost, positron emission tomography detector with high spatial resolution and DOI and TOF performance for PET/MRI.
  • An object of the present invention is to provide a photoelectric converter having a higher spatial resolution, comprising a silicon photomultiplier array, wherein the silicon photomultiplier array is spliced in a horizontal plane by i ⁇ j silicon photomultiplier tubes, i is the number of silicon photomultiplier tubes in the longitudinal direction on the horizontal plane, and j is the number of silicon photomultiplier tubes in the width direction on the horizontal plane, and both i and j are integers greater than or equal to 2.
  • Another object of the present invention is to provide a detector having the photoelectric converter, the detector comprising a scintillation crystal, an electronic system, the detector having the photoelectric converter, the scintillation crystal, the light guide, the silicon photoelectric
  • the multiplier array is sequentially coupled in sequence by an optical coupling agent.
  • Another object of the present invention is to provide a scanning device having the detector, the scanning device comprising a detecting device and a frame, the detecting device being mounted on the frame, the detecting device comprising the detector.
  • the present invention mainly adopts the photoelectric detection scheme of the silicon photomultiplier tube, because the silicon photomultiplier tube is small in size and compact in arrangement, and the appropriate size and number of silicon photomultiplier tubes are matched with the light guide of a suitable shape.
  • the photons of the scintillation crystal are diffused in the light guide, and the spatial information can be used to build a high spatial resolution PET detector, which ultimately improves the spatial resolution of the entire PET system, and is suitable for constructing PET detectors with DOI and TOF performance. Can be used for PET/MRI and low cost.
  • Photoelectric detection system scheme using silicon photomultiplier tube compared with the traditional large-volume photomultiplier tube scheme, the silicon photomultiplier tube is compact and compact, and the appropriate size and number of silicon photomultiplier tubes are matched with the shape of the light guide. After that, PET detectors with inherently high spatial resolution can be built to ultimately improve the spatial resolution of the entire PET system.
  • the silicon photomultiplier tube has a gain comparable to that of the photomultiplier tube (10 6 ).
  • the avalanche photodiode has a gain of only 10 4 and a large noise, which is composed of an avalanche photodiode.
  • PET detectors have lower performance.
  • silicon photomultiplier tube is small in size, it is very convenient to build a variety of detector structures, and is very suitable for building PET detectors with DOI detection capability, greatly improving the structure of PET detectors compared with the larger photomultiplier tubes. flexibility.
  • silicon photomultiplier tube has good time performance, can build TOF-capable PET detector to improve PET image quality.
  • the photomultiplier tube widely used in traditional PET detectors can not work normally in the magnetic field, which makes it difficult to develop photoelectric conversion devices in the PET/MRI dual-mode imaging system.
  • the silicon photomultiplier tube has high gain and the working state is not affected by the magnetic field. The best choice for core optoelectronic conversion devices in the PET/MRI solution.
  • the photomultiplier tube especially the position-sensitive photomultiplier tube, is too complicated due to the complicated production process, and the cost of the silicon photomultiplier tube is low when it is mass-produced, which can greatly reduce the cost of constructing the PET system.
  • Using three electronic pre-processing circuits can effectively reduce the number of channels of PET detectors built on silicon photomultiplier tubes without losing the position, energy and time information of gamma photon deposition, so that silicon-based photomultiplier tubes can be built. PET detectors are easier to implement.
  • FIG. 1 is a perspective view of a first embodiment of a photoelectric converter according to the present invention.
  • FIG. 2 is a perspective view of a second embodiment of the photoelectric converter of the present invention.
  • FIG. 3 is a perspective view of a third embodiment of the photoelectric converter of the present invention.
  • FIG. 4 is a schematic diagram of how the four-channel silicon photomultiplier signal of the present invention finally generates a 4-way weighted signal through a resistor network;
  • FIG. 5 is a schematic diagram of how the 16-channel silicon photomultiplier signal of the present invention finally generates a 4-way weighted signal through a more simple and clear resistor network;
  • FIG. 6 is a schematic diagram of how the 16-channel silicon photomultiplier signal of the present invention finally produces an 8-way weighted signal by simultaneously taking the cathode and anode signals of the silicon photomultiplier tube.
  • the silicon photomultiplier tube has the same gain as the photomultiplier tube, and has the advantages of small volume of the avalanche photodiode and insensitivity to the magnetic field, and has good time performance and low price in mass production.
  • the silicon photomultiplier tube combines the advantages of photomultiplier tube and avalanche photodiode, making full use of these advantages, and matching the corresponding light guide design, so that the photons of the scintillation crystal are diffused in the light guide, using this rich spatial distribution information to match the back end
  • the electronic design and position reading algorithm can obtain the gamma photon deposition position information more accurately, and make the spatial resolution of the PET detector break through the size limit of the silicon photomultiplier tube.
  • the present invention discloses a photoelectric converter with higher spatial resolution, DOI and TOF performance, can be used for PET/MRI, low cost, a detector having the same, and a detector having the same Scan the device.
  • the scanning device includes a detecting device and a frame, the detecting device is mounted on the frame, and the detecting device includes a detector.
  • the detector includes a scintillation crystal, an electronics system, a photoelectric converter, the optoelectronic converter comprising a silicon photomultiplier array and a lightguide coupled to the silicon photomultiplier array, the scintillation crystal, lightguide, silicon photomultiplier array Coupling is performed sequentially through the optical coupling agent in sequence.
  • the silicon photomultiplier tube is small in size, it is convenient to build a variety of detector structures, and is very suitable for building PET detectors with DOI detection capability. Compared with the larger photomultiplier tube, the structure of the PET detector is greatly improved. Flexibility, and the silicon photomultiplier tube has good time performance, can build TOF-capable PET detector to improve PET image quality.
  • the silicon photomultiplier tube array is spliced in a horizontal plane by i ⁇ j silicon photomultiplier tubes, and both i and j are integers greater than or equal to 2, and the scintillation crystal, light guide, and silicon photomultiplier tube array are sequentially pressed.
  • the sequence is coupled by an optical coupling agent.
  • the scintillation photons from the scintillation crystal pass through the light guide, light diffusion occurs.
  • the gamma photons can be accurately deposited in the scintillation crystal. position.
  • the silicon photomultiplier tube is small in size, low in price and tightly arranged.
  • the spatial resolution of the detector can be improved. The rate can ultimately improve the imaging quality of the system; at the same time, the silicon photomultiplier tube has a gain comparable to the gain of the photomultiplier tube (10 6 ), compared to the conventional photoelectric conversion device avalanche photodiode, its gain is only 10 4 , noise Larger, the PET detector consisting of avalanche photodiodes has lower performance.
  • the detector includes a scintillation crystal 1 for converting gamma photons into scintillation photons, a photoelectric converter 2 for converting scintillation photons into electrical pulse signals, and a deposition position, energy, and gamma photons calculated from electric pulse signals.
  • the photoelectric converter 2 includes a light guide 4 and a silicon photomultiplier array 5 coupled to the light guide 4, and the scintillation crystal 1, the light guide 4, and the silicon photomultiplier tube array 5 are sequentially coupled by an optical coupling agent in order.
  • the optical coupling agent can be an optical glue.
  • the light guide is an optical fiber or a fully cut transparent element or a non-fully cut transparent element or a continuous transparent element.
  • the material of the transparent member is ordinary inorganic glass or organic glass or scintillation crystal.
  • the scintillation crystal includes an array crystal composed of a single crystal strip, or an array crystal that is not completely cut, or an uncut continuous crystal, or a multilayer crystal in which a continuous crystal is combined with an array crystal.
  • the scintillation crystal is an inorganic scintillation crystal, and the material thereof is barium strontium silicate, barium silicate, barium bromide, barium silicate, barium silicate, barium fluoride, sodium iodide, barium iodide.
  • silicon photomultiplier tubes produced by their own companies, and generally have the following names: silicon photomultiplier (SiPM), multi-pixel photon counter (multi-pixel) Photon counter, MPPC), Geiger-mode avalanche photodiode (G-APD), digital silicon photomultiplier (digital silicon photomultiplier, dSiPM), although differently called, actually they all refer to the silicon photomultiplier tube of the present invention, and the principle functions are the same.
  • the silicon photomultiplier tube described in the present invention is only a general term, and the scope of protection of the present invention does not have different protection scopes due to different names, that is, others cannot subjectively think that changing the naming is different from the invention. .
  • the scope of actual protection of the silicon photomultiplier tube disclosed by the present invention includes products specified by different manufacturers in the prior art for different names of silicon photomultiplier tubes.
  • the single silicon photomultiplier tube has a detection area of between 1 x 1 mm 2 and 6 x 6 mm 2 and a micro pixel unit area of between 25 x 25 um 2 and 100 x 100 um 2 .
  • the shape of the light guide includes a cone, a cylinder, a rectangular parallelepiped, a cube, and a cone-like polyhedron. Of course, the light guide may have other shapes, which are not enumerated here.
  • the light guide comprises an optical fiber, a completely cut transparent element, an incompletely cut transparent element, a continuous transparent element or other transparent element, the material of which comprises ordinary inorganic glass, plexiglass, scintillation crystal.
  • the light guide comprises a P layer, the P range is between 0 and 4 layers, and all light guides add up to a thickness between 0.1 mm and 50 mm. As shown in FIGS. 1 to 2, only the case where the number m of photoconductive layers is equal to 1 is shown, which are continuous transparent elements.
  • the photoelectric converter comprises a silicon photomultiplier tube array formed by splicing i ⁇ j silicon photomultiplier tubes on a horizontal surface, wherein i is the number of silicon photomultiplier tubes in the longitudinal direction on the horizontal plane, and i is an integer greater than or equal to 2, j The number of silicon photomultiplier tubes in the width direction on the horizontal plane, j is an integer greater than or equal to 2; for a gamma photon deposition event, the photoelectric converter will generate k electrical pulse signals, where k is an integer greater than or equal to 4. .
  • the electronic system obtains gamma photon energy, position, and time information by processing a k-way electrical pulse signal.
  • the electronic system for processing the electric signal to obtain the position, energy and time information of the gamma photon is not preprocessed for the k electrical pulse signal, and directly reads the k electrical pulse signal one-to-one, using the maximum likelihood estimation method,
  • the artificial neural network localization algorithm calculates the gamma photon deposition location.
  • the electronic system for processing the electrical signal to obtain the position, energy and time information of the gamma photon can also preprocess the k electrical pulse signal to reduce the number of electronic channels.
  • the preprocessing circuit includes: Anger circuit, discrete proportional (discretized proportional) Counter, DPC) circuit, cross-wire circuit, obtains m electrical pulse signal, m is an integer greater than or equal to 4, less than or equal to k. As shown in FIG. 4 to FIG. 6, only the Anger circuit and the DPC circuit are respectively shown in the figure. The case of a cross-wire circuit.
  • the use of three electronic pre-processing circuits can effectively reduce the number of channels of PET detectors built on silicon photomultiplier tubes without losing the position, energy and time information of gamma photon deposition, making PET based on silicon photomultiplier tube construction.
  • the detector is easier to implement in engineering.
  • the electronic system uses a position algorithm to calculate a gamma photon deposition position based on m electrical pulse signals, and the position algorithm includes a center of gravity method, an Anger-Logic method, a maximum likelihood estimation method, and an artificial neural network localization algorithm.
  • the electronic system may also calculate a gamma photon deposition time based on m electrical pulse signals by using a time algorithm, which adds or equalizes m electrical pulse signals to obtain an additive electrical pulse signal, and extracts The time information of the summed electrical pulse signal is used as the deposition time of the gamma photon.
  • the specific method for obtaining the time includes a constant fraction discrimination (CFD) method, a leading edge discrimination (LED) method, a multi-voltage threshold (MVT) method, and a cycle time sampling ( Regular-time sampling (RTS) method.
  • CFD constant fraction discrimination
  • LED leading edge discrimination
  • MVT multi-voltage threshold
  • RTS Regular-time sampling
  • Embodiment 1 is a diagrammatic representation of Embodiment 1:
  • the detector includes a scintillation crystal 1, a photoelectric converter 2, and an electronic system 3.
  • the scintillation crystal 1 is composed of 12 ⁇ 12 scintillation crystal strips of the same size and spliced in a horizontal plane.
  • the bottom surface of the scintillation crystal 1 is directly coupled to the top surface of the light guide 4.
  • the silicon photomultiplier tube array 5 is composed of 4 x 4 silicon photomultiplier tubes of the same size, wherein the light guide has only one layer, the outer shape is a rectangular parallelepiped, the material is glass, and the thickness is 13 mm.
  • the DPC circuit is first used to reduce the channel number from 16 to 4, and then the Anger-Logic algorithm is used to obtain the gamma photon deposition position; the method for obtaining the gamma photon deposition time information is to add the 4 signals outputted by the DPC circuit. And, the time for extracting the sum signal is the deposition time of the gamma photon.
  • Embodiment 2 is a diagrammatic representation of Embodiment 1:
  • the detector includes a scintillation crystal 1, a photoelectric converter 2, and an electronic system 3.
  • the scintillation crystal 1 is composed of 12 ⁇ 12 scintillation crystal strips of the same size and spliced in a horizontal plane.
  • the bottom surface of the scintillation crystal 1 is directly coupled to the top surface of the light guide 4, and the silicon photomultiplier tube array 5 is composed of 4 ⁇ 4 silicon photomultiplier tubes of the same size, wherein the light guide is only one layer, the shape is a cone-like hexahedron, and the upper and lower surfaces are They are all square, and the four sides are trapezoidal.
  • the square surfaces of the upper and lower surfaces are not the same, and the thickness is 13mm.
  • the cross-wire circuit is first used to reduce the number of channels from 16 to 8, and then the maximum likelihood estimation method is used to obtain the gamma photon deposition position; the method for obtaining the gamma photon deposition time information is to output the cross-wire circuit.
  • the sum of the four signals in one direction is extracted, and the time for extracting the sum signal is the deposition time of the gamma photons.
  • Embodiment 3 is a diagrammatic representation of Embodiment 3
  • the position sensitive PET detector based on the silicon photomultiplier tube includes a scintillation crystal 1, a photoelectric converter 2, and an electronic system 3.
  • the scintillation crystal 1 is composed of 12 ⁇ 12 scintillation crystal strips of the same size and spliced in a horizontal plane.
  • the bottom surface of the scintillation crystal 1 is directly coupled to the top surface of the light guide 4, and the silicon photomultiplier tube array 5 is composed of 4 ⁇ 4 silicon photomultiplier tubes of the same size, wherein the light guide is only one layer, and the outer shape is a cone-like decahedron, that is, one
  • the cuboid is added with a hexahedron (the middle is a continuous light guide), and the upper and lower surfaces of the entire decahedron are square, the surface area is not the same, and the sides are four rectangles plus four trapezoids, and the total thickness is 13 mm.
  • the cross-wire circuit is first used to reduce the number of channels from 16 to 8, and then the maximum likelihood estimation method is used to obtain the gamma photon deposition position; the method for obtaining the gamma photon deposition time information is to output the cross-wire circuit.
  • the sum of the four signals in one direction is extracted, and the time for extracting the sum signal is the deposition time of the gamma photons.
  • FIG. 4 Anger circuit diagram, here shows how the four-channel silicon photomultiplier signal finally generates four-way weighted signals through the resistor network, and each channel signal is matched by different resistance values in each direction. The weighting is performed, and then the position of the silicon photomultiplier signal generation is calculated by the centroid method. According to this principle, it is possible to expand to 16, 16 channels and finally generate 4-way weighted new numbers without paying creative labor.
  • DPC circuit diagram here shows how the 16-channel silicon photomultiplier signal passes through a simpler and clearer resistor network to finally generate four-way weighted signals.
  • each way is different.
  • Silicon photomultiplier signals have different weights and then utilized
  • the Anger-Logic algorithm generates the position.
  • the resistance value in the circuit is only an indication. According to this principle, only a part of the resistance value or the number of the resistance can be changed without any creative work, and the silicon photomultiplier tube can be extended to other x ⁇ y channels. How the channel signal finally produces a 4-way weighted new number, x and y are integers greater than or equal to 2.
  • cross-wire circuit diagram here shows how the 16-channel silicon photomultiplier signal passes through the cathode and anode signals of the simultaneous desiliconization photomultiplier tube. Finally, the 8-way weighting signal is generated.
  • the circuit can The x ⁇ y-way silicon photomultiplier signal is reduced to x+y, x and y are integers greater than or equal to 2, and then the position is generated using maximum likelihood estimation or an artificial neural network algorithm.
  • it can be extended to other x ⁇ y-way silicon photomultiplier tubes without any creative labor, and both x and y are integers greater than or equal to two.
  • Photoelectric detection system scheme using silicon photomultiplier tube compared with the traditional large-volume photomultiplier tube scheme, the silicon photomultiplier tube is compact and compact, and the appropriate size and number of silicon photomultiplier tubes are matched with the shape of the light guide. After that, PET detectors with inherently high spatial resolution can be built to ultimately improve the spatial resolution of the entire PET system.
  • the silicon photomultiplier tube has a gain comparable to that of the photomultiplier tube (10 6 ).
  • the avalanche photodiode has a gain of only 10 4 and a large noise, which is composed of an avalanche photodiode.
  • PET detectors have lower performance.
  • Silicon photomultiplier tube is small in size, it is convenient to build a variety of detector structures, and is very suitable for building PET detectors with DOI detection capability. Compared with the larger photomultiplier tube, the structure of PET detector is greatly improved. flexibility.
  • Silicon photomultiplier tube has good time performance, can build TOF-capable PET detector to improve PET image quality.
  • the photomultiplier tube widely used in traditional PET detectors can not work normally in the magnetic field, which makes it difficult to develop photoelectric conversion devices in the PET/MRI dual-mode imaging system.
  • the silicon photomultiplier tube has high gain and the working state is not affected by the magnetic field. The best choice for core optoelectronic conversion devices in the PET/MRI solution.
  • Using three electronic pre-processing circuits can effectively reduce the number of channels of PET detectors built on silicon photomultiplier tubes without losing the position, energy and time information of gamma photon deposition, so that silicon-based photomultiplier tubes can be built. PET detectors are easier to implement.

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Abstract

一种光电转换器，其包括硅光电倍增管阵列（5）及与硅光电倍增管阵列（5）耦合的光导（4），硅光电倍增管阵列（5）为i×j个硅光电倍增管在水平面上拼接而成，i和j都是大于或等于2的整数。一种探测器，其包括闪烁晶体（1）、电子学系统（3）、光导（4）、硅光电倍增管。一种扫描设备，其包括探测装置和机架，探测装置包括探测器，探测器包括所述的光电转换器。硅光电倍增管体积小巧，排列紧实，适当尺寸与个数的硅光电倍增管搭配适合形状的光导后，可搭建高空间分辨率的PET探测器，提高整个PET系统的空间分辨率，并且适合搭建具有DOI和TOF性能的PET探测器，能用于PET/MRI、成本低廉。

Description

光电转换器、探测器及扫描设备
本申请要求于2014年11月14日提交中国专利局、申请号为201410648328.4、发明名称为“光电转换器、探测器及扫描设备”的中国专利申请的优先权，其全部内容通过引用结合在本申请中。
技术领域
本发明涉及核医学成像技术领域，特别是涉及一种光电转换器、具有该光电转换器的探测器及具有该探测器的扫描设备。
背景技术
正电子发射断层成像(Positron Emission Tomography，以下简称PET)探测器是PET成像设备中的关键装置，其主要功能是获得PET系统中γ光子沉积时的位置、时间和能量信息。PET探测器的性能直接决定了整个PET成像系统的性能，为了提高系统成像性能，希望所使用的正电子发射断层成像探测器具有高空间分辨率、好的时间分辨率、好的能量分辨率、高计数率等特性。
PET探测器中光电转换器的功能是将前端闪烁晶体输出的闪烁光子转化为对应的电脉冲，并进行倍增放大得到电子学系统可以处理的电脉冲信号。现有的PET探测器，其中光电转换器件通常采用光电倍增管、雪崩光电二极管、位置敏感型光电倍增管。
光电倍增管一般通过光电面将闪烁光子转化为光电子，然后通过多个倍增极对光电子进行倍增，最后从阳极输出电脉冲信号，光电倍增管增益一般为106左右。光电倍增管具有高增益、低噪声、快的时间响应等优点，这造就了大部分的临床PET光电转换器件均选择了光电倍增管。但其体积一般较大，这可能限制PET探测器的空间分辨率以及PET系统结构设计的灵活性；光电倍增管无法在磁场中正常工作，难以作为PET/MRI双模成像系统的光电转换器件。
雪崩光电二极管首先通过光阴极将闪烁光子转化为光电子，利用光电二极管的雪崩效应来对光电子进行倍增得到电脉冲信号，雪崩光电二极管可以在磁场中正常工作，在PET/MRI双模成像中系统表现出一定的潜力。雪崩光电二极管体积较小，可以用来设计空间分辨率更好的PET探测器，已经有人通过 雪崩光电二极管来设计对空间分辨率要求更高的小动物PET(M é lanie Bergeron，Jules Cadorette,，Jean-
Figure PCTCN2015092833-appb-000001
Beaudoin，Martin D.Lepage，Ghislain Robert，Vitali Selivanov，Marc-Andr é T é trault，Nicolas Viscogliosi，Jeffrey P.Norenberg，R é jean Fontaine，and Roger Lecomte，“Performance Evaluation of the LabPET APD-Based Digital PET Scanner，”IEEE TRANSACTIONS ON NUCLEAR SCIENCE,VOL.56,NO.1,FEBRUARY 2009)。但是雪崩光电二极管有着天然的缺陷，增益不够高，大概104，噪声较大，会影响PET探测器的性能。
位置敏感型光电倍增管具备所有光电倍增管的优点，而且能够实现较高的空间分辨率，有研究组已经通过位置敏感型光电倍增管实现了对空间分辨率要求更高的小动物PET(Qingguo Xie,Yuanbao Chen,Jun Zhu,Jingjing Liu,Xi Wang,Xin Chen,Ming Niu,Zhongyi Wu,Daoming Xi,Luyao Wang,Peng Xiao,Chin-Tu Chen,Chien-Min Kao“Implementation of LYSO-PSPMT Block Detector with an All-Digital DAQ System,”in IEEE Transactions on Nuclear Science,pp.1487-1494,2013)，并达到了较好的系统性能。但是位置敏感型光电倍增管价格十分昂贵，会增加PET系统的成本。
近年来，一种成本低廉的硅光电倍增管引起了人们的注意，硅光电倍增管由运行在盖革模式下的雪崩光电二极管微像素单元阵列组成，增益为106，该增益与光电倍增管媲美，噪声较低，体积小，对磁场不敏感，且其时间性能良好，大量生产时价格低廉，适合用来搭建PET探测器(Qingguo Xie,Robert G.Wagner,Gary Drake,Patrick DeLurgio,Yun Dong,Chin-Tu Chen,Chien-Min Kao,“Performance Evaluation of Multi-Pixel Photon Counters for PET Imaging,”in Conference Record of the 2007 IEEE Nuclear Science Symposium,vol.2,pp.969-974,2007)。采用硅光电倍增管作为光电转换器件，对比传统体积较大光电倍增管，硅光电倍增管体积小巧，排列紧实，适合搭建高空间分辨率PET探测器，最终提高整个PET系统的空间分辨率；硅光电倍增管体积小巧，很方便搭建多种探测器结构，并且很适合搭建具有沉积深度(Depth of Interaction，简称DOI)探测能力的PET探测器，相对于体积较大的光电倍增管，大大提 高了PET系统结构搭建的灵活性；硅光电倍增管具有良好的时间性能，可以搭建具有飞行时间(Time-of-Flight，简称TOF)能力的PET探测器，提高PET图像质量；硅光电倍增管增益高，工作状态不受磁场影响，是PET/MRI方案中核心光电转换器件的最佳选择。光电倍增管尤其是位置敏感型光电倍增管成本一直居高不下，硅光电倍增管大量生产时价格低廉，可大幅度降低PET系统的成本。
硅光电倍增管具有与光电倍增管媲美的增益，又具有雪崩光电二极管体积小、对磁场的不敏感的优点，且其时间性能良好。硅光电倍增管兼具了光电倍增管和雪崩光电二极管的优点，如果正电子发射断层成像探测器利用硅光电倍增管作为光电转换器件并能够较好的实现位置分辨，便可以基于硅光电倍增管设计并搭建空间分辨率更高的、具有DOI和TOF性能的、能用于PET/MRI的、成本低廉的正电子发射断层成像探测器。
因此，有必要提供一种新型的应用硅光电倍增管的光电转换器，以克服现有技术中探测器的缺陷。
发明内容
有鉴于此，本发明的目的在于提供一种空间分辨率更高的光电转换器、具有该光电转换器的探测器及具有该探测器的扫描设备。
本发明一个目的在于提供一种空间分辨率更高的光电转换器，其包括硅光电倍增管阵列，所述硅光电倍增管阵列为i×j个硅光电倍增管在水平面上拼接成，所述i为水平面上长度方向的硅光电倍增管个数，所述j为水平面上宽度方向的硅光电倍增管个数，所述i和j都是大于或等于2的整数。
本发明另一个目的在于提供一种具有该光电转换器的探测器，该探测器包括闪烁晶体、电子学系统，所述探测器具有所述的光电转换器，所述闪烁晶体、光导、硅光电倍增管阵列依次按顺序通过光学耦合剂进行耦合。
本发明另一个目的在于提供一种具有该探测器的扫描设备，该扫描设备包括探测装置和机架，所述探测装置安装于所述机架上，所述探测装置包括所述的探测器。
从上述技术方案可以看出，本发明主要采用硅光电倍增管的光电探测方案，因为硅光电倍增管体积小巧，排列紧实，适当尺寸与个数的硅光电倍增管搭配适合形状的光导后，使闪烁晶体的光子在光导产生扩散，利用丰富的空间信息，可搭建高空间分辨率PET探测器，最终提高整个PET系统的空间分辨率，并且很适合搭建具有DOI和TOF性能的PET探测器，能用于PET/MRI、成本低廉。
与现有技术相比，本发明的有益效果是：
1、采用硅光电倍增管的光电探测系统方案，对比传统较大体积光电倍增管的方案，硅光电倍增管体积小巧，排列紧实，适当尺寸与个数的硅光电倍增管搭配适合形状的光导后，可搭建固有空间分辨率很高的PET探测器，最终提高整个PET系统的空间分辨率。
2、硅光电倍增管具有与光电倍增管媲美的增益(106)可以接受的噪声，对比传统的光电转换器件雪崩光二极管，其增益只有104，噪声较大，使得由雪崩光电二极管组成的PET探测器性能较低。
3、硅光电倍增管体积小巧，很方便搭建多种探测器结构，并且很适合搭建具有DOI探测能力的PET探测器，相对于体积较大的光电倍增管，大大提高了PET探测器结构搭建的灵活性。
4、硅光电倍增管具有良好的时间性能，可以搭建具有TOF能力的PET探测器，提高PET图像质量。
5、传统PET探测器中广泛使用的光电倍增管无法在磁场中正常工作，使得研发PET/MRI双模成像系统存在光电转换器件困难，硅光电倍增管增益高，工作状态不受磁场影响，是PET/MRI方案中核心光电转换器件的最佳选择。
6、光电倍增管尤其是位置敏感型光电倍增管由于生产过程过于复杂，成本一直居高不下，硅光电倍增管大量生产时价格低廉，可大幅度降低搭建PET系统的成本。
7、采用三种电子学预处理电路既可以有效的减少基于硅光电倍增管搭建的PET探测器的通道数，又不丢失γ光子沉积的位置、能量、时间信息，使得基于硅光电倍增管搭建的PET探测器更易于工程实现。
附图说明
图1为本发明光电转换器第一实施例的立体示意图；
图2为本发明光电转换器第二实施例的立体示意图；
图3为本发明光电转换器第三实施例的立体示意图；
图4为本发明4个通道的硅光电倍增管信号通过电阻网络如何最后产生4路加权信号的示意图；
图5为本发明16个通道的硅光电倍增管信号通过更加简单明了的电阻网络如何最后产生4路加权信号的示意图；
图6为本发明16个通道的硅光电倍增管信号通过同时取硅光电倍增管的阴极和阳极信号如何最后产生8路加权信号的示意图。
具体实施方式
硅光电倍增管具有与光电倍增管媲美的增益，又具有雪崩光电二极管体积小、对磁场的不敏感的优点，且其时间性能良好、大量生产时价格低廉。硅光电倍增管兼具了光电倍增管和雪崩光电二极管的优点，充分利用这些优点，配合相应的光导设计，使闪烁晶体的光子在光导中产生扩散，利用这丰富的空间分布信息，配合后端的电子学设计以及位置读出算法，可以更加准确的获得γ光子沉积位置信息，使PET探测器空间分辨率突破硅光电倍增管尺寸限制。
基于以上分析，本发明公开了一种空间分辨率更高、具有DOI和TOF性能、能用于PET/MRI、成本低廉的光电转换器、具有该光电转换器的探测器及具有该探测器的扫描设备。
所述扫描设备包括探测装置和机架，所述探测装置安装于所述机架上，所述探测装置包括探测器。
所述探测器包括闪烁晶体、电子学系统、光电转换器，所述光电转换器包括硅光电倍增管阵列及与硅光电倍增管阵列耦合的光导，所述闪烁晶体、光导、硅光电倍增管阵列依次按顺序通过光学耦合剂进行耦合。硅光电倍增管体积小巧，很方便搭建多种探测器结构，并且很适合搭建具有DOI探测能力的PET探测器，相对于体积较大的光电倍增管，大大提高了PET探测器结构搭建的 灵活性，且硅光电倍增管具有良好的时间性能，可以搭建具有TOF能力的PET探测器，提高PET图像质量。
所述硅光电倍增管阵列为i×j个硅光电倍增管在水平面上拼接成，所述i和j都是大于或等于2的整数，所述闪烁晶体、光导、硅光电倍增管阵列依次按顺序通过光学耦合剂进行耦合。来自于闪烁晶体的闪烁光子经由光导时，发生光扩散，根据耦合在光导下各个硅光电倍增管收到的数目的光子的分布，结合位置算法，可以准确的获得γ光子在闪烁晶体中沉积的位置。通过采用硅光电倍增管的光电探测系统方案，对比传统较大体积光电倍增管的方案，硅光电倍增管体积小巧，价格低廉、排列紧实，搭建PET探测器时，可以提高探测器的空间分辨率，最终可以提高系统的成像质量；同时，硅光电倍增管具有与光电倍增管媲美的增益(106)可以接受的噪声，对比传统的光电转换器件雪崩光二极管，其增益只有104，噪声较大，使得由雪崩光电二极管组成的PET探测器性能较低。
如图1所示，该探测器包含将γ光子转化为闪烁光子的闪烁晶体1、将闪烁光子转换为电脉冲信号的光电转换器2以及根据电脉冲信号来计算γ光子的沉积位置、能量、时间信号的电子学系统3。其中，光电转换器2包括光导4及与光导4耦合硅光电倍增管阵列5，闪烁晶体1、光导4、硅光电倍增管阵列5依次按顺序通过光学耦合剂进行耦合。光学耦合剂可以是光学胶水。
所述光导为光纤或完全切割的透明元件或不完全切割的透明元件或连续的透明元件。所述透明元件的材料为普通无机玻璃或有机玻璃或闪烁晶体。
所述闪烁晶体包括单个晶体条组成的阵列晶体，或者未完全切割的阵列晶体，或者未切割的连续晶体，或者连续晶体与阵列晶体组合的多层晶体。
所述闪烁晶体为无机闪烁晶体，其材料为锗酸铋、硅酸镥、溴化镧、硅酸钇镥、硅酸钇、氟化钡、碘化钠、碘化铯。
现有技术中，不同厂商对自己公司生产的硅光电倍增管都有各自不同的命名，大致有如下几种叫法：硅光电倍增器(silicon photomultiplier，SiPM)、多像素光子计数器(multi-pixel photon counter，MPPC)、盖革模式雪崩光电二极管(Geiger-mode avalanche photodiode，G-APD)、数字硅光电倍增器 (digital silicon photomultiplier，dSiPM)，虽然叫法不同，但是实际上他们均是指本发明所述的硅光电倍增管，其原理功能均相同。本发明描述的硅光电倍增管只是一个统称，本发明的保护范围并不会因为叫法的不同而出现不同的保护范围，即他人不可以主观上认为改变命名就与本发明属于不同的发明创造。本发明所揭示的硅光电倍增管的实际保护的范围囊括现有技术中各大厂商对硅光电倍增管的不同叫法所指定的产品。
所述单颗硅光电倍增管探测面积在1×1mm2至6×6mm2之间，微像素单元面积在25×25um2至100×100um2之间。
所述光导的形状包括圆锥体、圆柱体，长方体、正方体及类锥形多面体，当然光导也可以是其他形状，在此不再一一进行列举。所述光导包括光纤、完全切割的透明元件、不完全切割的透明元件、连续的透明元件或其他透明元件，所述透明元件的材料包括普通无机玻璃、有机玻璃、闪烁晶体。所述的光导包括P层，所述P的范围在0至4层之间，所有光导加起来的厚度在0.1mm至50mm之间。如图1至图2所示，图中仅显示了光导层数m均等于1，均为连续的透明元件的情况。
所述光电转换器包括i×j硅光电倍增管在水平面上拼接成的硅光电倍增管阵列，其中i为水平面上长度方向的硅光电倍增管个数，i为大于或等于2的整数，j为水平面上宽度方向的硅光电倍增管个数，j为大于或等于2的整数；对于一次γ光子沉积事件，光电转换器将会产生k路电脉冲信号，其中k为大于或等于4的整数。所述电子学系统通过处理k路电脉冲信号获得γ光子能量、位置、时间信息。所述处理电信号得到γ光子的位置、能量及时间信息的电子学系统对于k路电脉冲信号不做预处理，直接对k路电脉冲信号一对一读出，使用最大似然估计法、人工神经网络定位算法计算γ光子沉积位置。
所述处理电信号得到γ光子的位置、能量及时间信息的电子学系统对于k路电脉冲信号也可以进行预处理，减少电子学通道数，预处理电路包括：Anger电路，离散正比(discretized proportional counter，DPC)电路，十字交叉(cross-wire)电路，得到m路电脉冲信号，m为大于或等于4，小于或等于k的整数。如图4至图6所示，图中仅分别显示了Anger电路，DPC电路， cross-wire电路的情况。采用三种电子学预处理电路既可以有效的减少基于硅光电倍增管搭建的PET探测器的通道数，又不丢失γ光子沉积的位置、能量、时间信息，使得基于硅光电倍增管搭建的PET探测器更易于工程实现。
所述电子学系统采用位置算法根据m个电脉冲信号计算出γ光子沉积位置，所述位置算法包括重心法、Anger-Logic法、最大似然估计法、人工神经网络定位算法。
所述电子学系统也可以采用时间算法根据m个电脉冲信号计算出γ光子沉积时间，所述时间算法为将m个电脉冲信号加和或者加权后加和获得一个加和电脉冲信号，提取加和电脉冲信号的时间信息作为γ光子的沉积时间。
所述求取时间的具体方法包括恒比甄别器(constant fraction discrimination，CFD)方法、边沿触发(leading edge discrimination，LED)方法、多电压阈值(multi-voltage threshold，MVT)方法、周期时间采样(regular-time sampling，RTS)方法。
下面将结合本发明实施例中的附图，对本发明实施例中的技术方案进行详细地描述，显然，所描述的实施例仅仅是本发明一部分实施例，而不是全部的实施例。基于本发明中的实施例，本领域普通技术人员在没有做出创造性劳动的前提下所获得的所有其他实施例，都属于本发明保护的范围。
实施例一：
如图1所示：探测器包括闪烁晶体1、光电转换器2、电子学系统3。其中闪烁晶体1为由12×12个相同尺寸的闪烁晶体条在水平面拼接构成。闪烁晶体1的底面直接与光导4的顶面耦合，硅光电倍增管阵列5由4×4个相同尺寸的硅光电倍增管，其中光导仅有一层，外形为长方体，材料为玻璃，厚度13mm。电子学系统3中，首先采用DPC电路，将通道数从16降至4，然后采用Anger-Logic算法获得γ光子沉积位置；获取γ光子沉积时间信息的方法为将DPC电路输出的4路信号加和，提取加和信号的时间为γ光子的沉积时间。
实施例二：
如图2所示：探测器包括闪烁晶体1、光电转换器2、电子学系统3。其中闪烁晶体1为由12×12个相同尺寸的闪烁晶体条在水平面拼接构成。闪烁晶体1的底面直接与光导4的顶面耦合，硅光电倍增管阵列5由4×4个相同尺寸的硅光电倍增管，其中光导仅为一层，外形为类锥形六面体，上下面表面均为正方形，四个侧面为梯形，上下表面正方形面积并不相同，厚度13mm。电子学系统3中，首先采用cross-wire电路，将通道数从16降至8，然后采用最大似然估计法获得γ光子沉积位置；获取γ光子沉积时间信息的方法为将cross-wire电路输出的一个方向的4路信号加和，提取加和信号的时间为γ光子的沉积时间。
实施例三：
如图3所示：基于硅光电倍增管的位置敏感PET探测器包括闪烁晶体1、光电转换器2、电子学系统3。其中闪烁晶体1为由12×12个相同尺寸的闪烁晶体条在水平面拼接构成。闪烁晶体1的底面直接与光导4的顶面耦合，硅光电倍增管阵列5由4×4个相同尺寸的硅光电倍增管，其中光导仅为一层，外形为类锥形十面体，即一个长方体加上一个六面体(中间为连续光导)，整个十面体上下两个表面为正方形，表面面积并不相同，侧面是四个长方形加四个梯形，总厚度13mm。电子学系统3中，首先采用cross-wire电路，将通道数从16降至8，然后采用最大似然估计法获得γ光子沉积位置；获取γ光子沉积时间信息的方法为将cross-wire电路输出的一个方向的4路信号加和，提取加和信号的时间为γ光子的沉积时间。
如图4所示：Anger电路示意图，此处示意了4个通道的硅光电倍增管信号通过电阻网络如何最后产生4路加权信号，在每一个方向上通过不同电阻值配比对每个通道信号进行加权，然后利用重心法计算硅光电倍增管信号产生位置。按照该原理可以在不付出创造性劳动的情况下，扩展到16、64通道如何最后产生4路加权新号。
如图5所示：DPC电路示意图，此处示意了16个通道的硅光电倍增管信号通过更加简单明了的电阻网络如何最后产生4路加权信号，显然，在该电路中，每一路不同位置的硅光电倍增管信号都有不同的加权值，然后利用 Anger-Logic算法产生位置。显然，该电路中电阻值仅仅是示意，按照该原理可以在不付出创造性劳动的情况下，仅仅改变部分电阻值或改变电阻的个数，便可以扩展到其它x×y路的硅光电倍增管通道信号如何最后产生4路加权新号的情况，x和y均为大于或等于2的整数。
如图6所示：cross-wire电路示意图，此处示意了16个通道的硅光电倍增管信号通过同时去硅光电倍增管的阴极和阳极信号如何最后产生8路加权信号，显然，该电路可以使x×y路的硅光电倍增管信号减少到x+y路，x和y均为大于或等于2的整数，然后利用最大似然估计法或者人工神经网络算法产生位置。显然，按照该原理可以在不付出创造性劳动的情况下，便可以扩展到其它x×y路的硅光电倍增管的情况，x和y均为大于或等于2的整数。
综上所述，与现有技术相比，本发明的有益效果包括：
1、采用硅光电倍增管的光电探测系统方案，对比传统较大体积光电倍增管的方案，硅光电倍增管体积小巧，排列紧实，适当尺寸与个数的硅光电倍增管搭配适合形状的光导后，可搭建固有空间分辨率很高的PET探测器，最终提高整个PET系统的空间分辨率。
2.硅光电倍增管具有与光电倍增管媲美的增益(106)可以接受的噪声，对比传统的光电转换器件雪崩光二极管，其增益只有104，噪声较大，使得由雪崩光电二极管组成的PET探测器性能较低。
3.硅光电倍增管体积小巧，很方便搭建多种探测器结构，并且很适合搭建具有DOI探测能力的PET探测器，相对于体积较大的光电倍增管，大大提高了PET探测器结构搭建的灵活性。
4.硅光电倍增管具有良好的时间性能，可以搭建具有TOF能力的PET探测器，提高PET图像质量。
5.传统PET探测器中广泛使用的光电倍增管无法在磁场中正常工作，使得研发PET/MRI双模成像系统存在光电转换器件困难，硅光电倍增管增益高，工作状态不受磁场影响，是PET/MRI方案中核心光电转换器件的最佳选择。
6.光电倍增管由于生产过程过于复杂，成本一直居高不下，硅光电倍增管大量生产时价格低廉，可大幅度降低搭建PET系统的成本。
7、采用三种电子学预处理电路既可以有效的减少基于硅光电倍增管搭建的PET探测器的通道数，又不丢失γ光子沉积的位置、能量、时间信息，使得基于硅光电倍增管搭建的PET探测器更易于工程实现。

Claims (15)

  1. 一种光电转换器，其特征在于：所述光电转换器包括硅光电倍增管阵列，所述硅光电倍增管阵列为i×j个硅光电倍增管在水平面上拼接成，所述i为水平面上长度方向的硅光电倍增管个数，所述j为水平面上宽度方向的硅光电倍增管个数，所述i和j都是大于或等于2的整数。
  2. 根据权利要求1所述的光电转换器，其特征在于：所述光电转换器还包括与硅光电倍增管阵列耦合的光导。
  3. 根据权利要求2所述的光电转换器，其特征在于：所述光导的形状为圆锥体或圆柱体或长方体或正方体或类锥形多面体。
  4. 根据权利要求2所述的光电转换器，其特征在于：所述光导为光纤或完全切割的透明元件或不完全切割的透明元件或连续的透明元件。
  5. 根据权利要求4所述的光电转换器，其特征在于：所述透明元件的材料为普通无机玻璃或有机玻璃或闪烁晶体。
  6. 根据权利要求2所述的光电转换器，其特征在于：所述光导包括P层，所述P的范围在0至4层之间。
  7. 根据权利要求6所述的光电转换器，其特征在于：所有所述光导加起来的厚度范围在0.1mm至50mm之间。
  8. 一种探测器，其包括闪烁晶体、电子学系统，其特征在于：所述探测器具有权利要求1至7任一所述的光电转换器，所述闪烁晶体、光导、硅光电倍增管阵列依次按顺序通过光学耦合剂进行耦合。
  9. 根据权利要求8所述的探测器，其特征在于：所述闪烁晶体为单个晶体条组成的阵列晶体，或者未完全切割的阵列晶体，或者未切割的连续晶体，或者连续晶体与阵列晶体组合的多层晶体。
  10. 根据权利要求8所述的探测器，其特征在于：对于一次γ光子沉积事件，所述光电转换器将会产生k路电脉冲信号，其中k为大于或等于4的整数，所述电子学系统通过处理k路电脉冲信号获得γ光子能量、位置、时间信息。
  11. 根据权利要求10所述的探测器，其特征在于：所述电子学系统对于k路电脉冲信号不进行预处理，直接对k路电脉冲信号一对一读出，计算γ光子沉积位置。
  12. 根据权利要求10所述的探测器，其特征在于：所述电子学系统对于k路电脉冲信号进行预处理，减少电子学通道数，得到m路电脉冲信号，m为大于或等于4，小于或等于k的整数。
  13. 根据权利要求12所述的探测器，其特征在于：所述电子学系统采用位置算法根据m个电脉冲信号计算出γ光子沉积位置。
  14. 根据权利要求12所述的探测器，其特征在于：所述电子学系统采用时间算法根据m个电脉冲信号计算出γ光子沉积时间，所述时间算法为将m个电脉冲信号加和或者加权后加和获得一个加和电脉冲信号，提取加和电脉冲信号的时间信息作为γ光子的沉积时间。
  15. 一种扫描设备，其包括探测装置和机架，所述探测装置安装于所述机架上，所述探测装置包括探测器，其特征在于：所述探测器具有权利要求1至7任一所述的光电转换器。
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