Cadmium zinc telluride positron emission tomography system and signal correction algorithm

Abstract

The invention discloses a CdZnTe positron emission tomography imaging system and a signal correction algorithm. The imaging system forms a ring detector unit with a CdZnTe detection module to form a single-ring PET or a multi-ring PET. The CdZnTe detection module includes a CdZnTe crystal module, a charge-sensitive preamplifier module and a filter amplifier module. The CdZnTe crystal module includes a CdZnTe crystal and an electrode. The output signal of the crystal passes through the charge-sensitive preamplifier module, the filter amplifier module and the signal readout module in sequence. The PET imaging system of the invention has a compact structure and is flexible to use, and improves the detection efficiency and spatial resolution of the detector. The invention proposes a method for calculating the three-dimensional position of gamma rays, a method for correcting the time resolution and an interpolation algorithm for charge sharing events for the output signal of the CdZnTe detector, effectively solving the problem of poor time performance of the CdZnTe detector and the resolution loss caused by the charge sharing effect, and improving the time response, counting rate and spatial resolution of the PET system.

Description

Technical Field

[0001] The present invention belongs to the technical field of medical diagnostic imaging equipment, and in particular relates to a cadmium zinc telluride positron emission tomography system and a signal correction algorithm. Background Art

[0002] Positron emission tomography (PET) is one of the most advanced large-scale three-dimensional imaging technologies for medical diagnosis. This technology has become one of the important methods for tumor examination, cardiovascular and cerebrovascular diseases and neurological diseases. At present, the detectors used for PET are mainly LSO crystals, BGO crystals, NaI crystals, GSO crystals, etc., and their spatial resolution can only meet the various functional imaging needs of the human body, and cannot meet the requirements of high-precision imaging of small animals. At the same time, large-scale PET equipment is expensive and occupies a large space, which greatly limits the application of current PET imaging systems in scientific research or other related fields. Most of the currently commonly used small animal PET can only obtain 2-dimensional positions, and cannot take into account both spatial resolution and sensitivity. In addition, its structural flexibility is poor, and the use of photoelectric conversion devices and some analog signal processing modules increases the power consumption of the instrument. Summary of the invention

[0003] In view of the deficiencies pointed out in the above background technology, the present invention provides a cadmium zinc telluride positron emission tomography (CZT-PET) imaging system and a CdZnTe detector output signal correction algorithm. The CZT-PET imaging system has a smaller detector volume and can be closer to the lesion to achieve high sensitivity and high counting rate. The output signal of the CdZnTe detector can effectively solve the problem of poor time performance of the CdZnTe detector through the correction algorithm, thereby improving the time response, counting rate and spatial resolution of the PET system.

[0004] To achieve the above object, the technical solution adopted by the present invention is:

[0005] The cadmium zinc telluride (CZT) positron emission tomography (PET) imaging system is used for imaging small animals or human organs. It uses cadmium zinc telluride (CZT) semiconductor as the detector body to form an annular detector unit, forming a single-ring PET or multi-ring PET. The single-ring PET can perform full-range imaging on most experimental small animals; in the multi-ring PET, multiple annular detector units are arranged in sequence along the axial direction, and the axial field of view is increased. Without moving the target object, full-range imaging can be achieved at one time.

[0006] The ring detector unit is composed of several small CdZnTe detection modules, each of which has a complete preamplifier module, that is, each CdZnTe detection module includes a CdZnTe crystal module, a charge-sensitive preamplifier module and a filter amplifier module, and its probe part is composed of a whole CdZnTe crystal module. The CdZnTe crystal module includes a CdZnTe crystal and an electrode. The electrode type of the crystal adopts an orthogonal strip electrode structure or a pixel electrode structure; the output signal of the CdZnTe crystal is processed by a charge-sensitive preamplifier with high gain and low noise. The amplitude of the processed signal is still only tens of millivolts, which is a relatively small amplitude. The electrical signal is easy to cause signal accumulation, resulting in the distortion of the collected signal, causing a large number of useful nuclear signals to be lost and distorted, and its baseline is extremely unstable. A large baseline value will cause damage to the back-end data acquisition system. Therefore, the signal processed by the charge-sensitive preamplifier is connected to the filter amplifier circuit. The filter amplifier circuit adopts a two-stage amplifier circuit to amplify the preamplifier output signal, and retains the shorter leading edge time of the signal, and greatly reduces the longer falling edge time of the signal, as well as the baseline recovery function to ensure better matching with the back-end data acquisition system. The signal output by the filter amplifier module enters the signal readout module.

[0007] The radial and axial fields of view of the overall CZT-PET can flexibly add or reduce PET detector modules according to the size of the target object. For small animals, CZT-PET can use fewer detector modules, and its radial diameter can be set to about 10 cm, which can reduce the overall system cost and improve resolution. For human brain imaging, detectors can be flexibly added according to the shape of the object to be detected, and detection can be performed at different positions. The radial diameter can be set to about 30 cm. For whole-body imaging of the human body, the number of detector rings can be increased in the axial direction to improve the sensitivity of the system. The radial diameter can be set to about 70 cm.

[0008] Preferably, in the orthogonal strip electrode structure, 16 parallel strip electrodes are respectively deposited on two opposite sides of the CdZnTe crystal; in the pixel type electrode structure, 16×16 square electrodes are deposited on the anode on one side of the CdZnTe crystal, and a planar electrode is used as the cathode on the other side.

[0009] Preferably, in the CdZnTe crystal module, a plurality of charge-sensitive preamplifiers are integrated via an ASIC chip.

[0010] The electrodes in the CdZnTe crystal module are made of metal gold material, and the spacing between adjacent strip electrodes on the same side in the orthogonal strip electrode structure and the spacing between adjacent square electrodes in the pixel type electrode structure are both 0.1 mm.

[0011] In addition, an electrode protection ring is set around the electrode in the CdZnTe crystal module to effectively reduce the end face leakage current of the detector.

[0012] The signal readout method of a single CdZnTe PET detector is as follows: the multi-channel front-end signals of a single CdZnTe crystal are converted into four-channel signal outputs. For an orthogonal strip electrode structure, the X signal and Y signal of the CdZnTe crystal are each converted into two-channel signal outputs; for a pixel-type electrode structure, all pixel anode signals of the CdZnTe crystal are converted into four-channel signal outputs.

[0013] The probe module of a single CdZnTe detection module is composed of four CdZnTe crystals that are laid or stacked to form the probe module of the single CdZnTe detection module, and the output signals of the four CdZnTe crystal detectors are compressed twice into 4-channel signals as the output signals of the single CdZnTe detection module.

[0014] For a CdZnTe positron emission tomography system, in order to effectively solve the problem of poor time performance of the CdZnTe detector and resolution loss caused by the charge sharing effect, the present invention further provides a detector output signal correction algorithm for the CdZnTe positron emission tomography system, which performs correction calculations from three aspects, including a method for calculating the three-dimensional position of gamma rays, a method for correcting the time resolution, and an interpolation algorithm for charge sharing events.

[0015] 1. Calculation method of the three-dimensional position of gamma rays in CdZnTe PET detectors

[0016] 1. The method for calculating the three-dimensional position of gamma rays of the CdZnTe PET detector (CZT-C32) with orthogonal strip electrode structure is:

[0017] The three-dimensional position (x, y, z) of an interaction event in a CdZnTe crystal is calculated as follows:

[0018]

[0019]

[0020]

[0021] Where G represents the distance between the edge electrode and the edge of the CdZnTe crystal, n represents the maximum induced current at the nth electrode, P represents the width of the anode strip, and A n Indicates the signal peak value at the nth anode, C n represents the signal peak value at the nth cathode, and D represents the thickness of the CdZnTe crystal.

[0022] 2. The method for calculating the three-dimensional position of gamma rays of the CdZnTe PET detector (CZT-P256) with a pixel-type electrode structure is:

[0023] The three-dimensional position (x, y, z) of an interaction event in a CdZnTe crystal is calculated as follows:

[0024]

[0025]

[0026]

[0027] Where G represents the distance between the edge electrode and the edge of the CdZnTe crystal, i and j represent the maximum induced current on the pixel electrode at the i-th row and j-th column, P represents the width of the anode strip, and A ij represents the signal peak value at the anode in the i-th row and j-th column, C represents the signal peak value at the cathode, and D represents the thickness of the CdZnTe crystal.

[0028] 2. Correction Method for the Time Resolution of CdZnTe PET Detector

[0029] Since the time performance of CZT detector is low, the rising edge of its cathode signal is longer and the rising edge of the anode signal is shorter. For the anode signal, the same energy is deposited at different positions, and the signal start time is different due to the different drift time of the electron. Therefore, the traditional detector output signal time discrimination method will cause a large error. Based on the phenomenon that the drift time of the electrons in the full energy peak event and the cathode-anode ratio maintain a linear relationship at most depths, the present invention proposes a method for obtaining a more accurate gamma reaction time ΔT in the cadmium zinc telluride detector, that is, subtracting the electron drift time from the measurement time of the anode signal, and the relationship is as follows:

[0030] ΔT=T CZT -T drift , (3-1)

[0031] Among them, T CZT is the readout time of the gamma-ray photon in the CdZnTe detector, T drift is the electron drift time; this correction method can greatly improve the time resolution of the anode signal in the CdZnTe detector.

[0032] 3. Interpolation Algorithm for Charge Sharing Events in CdZnTe PET Detectors

[0033] Charge sharing events reduce the spatial resolution of the PET system during PET image reconstruction. To solve this problem, the present invention proposes the following interpolation algorithm:

[0034] The charge cloud is detected by multiple pixels. According to the relationship that the signal amplitude of adjacent anode pixels is proportional to the charge collected, the signal amplitudes of three adjacent and sequentially arranged anode pixels are A and B, respectively. 1 , A 2 , A 3 , calculate the ratio R of the signal amplitudes of two adjacent anode electrode pixels:

[0035] R 21 =A 2 / A 1 ; (4-1)

[0036] R 23 =A 2 / A 3 ; (4-2)

[0037] According to the ratio R 21 and R 23 , A 2 The corresponding anode pixel range is divided into five regions, which are respectively recorded as regions A, B, C, D, and E. The division ranges are as follows:

[0038] 1<R 21 <3, area A;

[0039] 3<R 21 <9, area B;

[0040] 9<R 21 or 9<R 23 , is region C;

[0041] 3<R 23 <9, it is area D;

[0042] 1<R 23 <3, it is area E;

[0043] The size of each divided area is one-fifth of the sum of the anode pixel and the anode gap, which is equivalent to using the interpolation algorithm of shared events to further reduce the anode pixel size to one-fifth, thereby determining the three-dimensional position of the charge sharing event and improving the position resolution of the charge sharing event. At the same time, the counting rates of each area are basically the same, thereby also improving the counting rate of the system.

[0044] Compared with the shortcomings and deficiencies of the prior art, the present invention has the following beneficial effects:

[0045] (1) The PET imaging system of the present invention uses a cadmium zinc telluride semiconductor detector, which has an excellent energy resolution of nearly 2% for 511keV annihilation gamma rays, which can greatly improve the accuracy of the ray energy selection range, thereby providing a strong guarantee for signal selection, facilitating the rejection of multiple scattering events and the selection of full-energy peak events, thereby achieving the purpose of improving the spatial resolution and signal-to-noise ratio of the system.

[0046] (2) The electrode of the CdZnTe semiconductor detector is made by the evaporation process. Compared with the current scintillator detector cutting process of the commercial PET system, the evaporation process is conducive to making the pixel reach the hundred-micron level, the cost is controllable, and the purity and uniformity of the evaporated electrode are good, which can provide the set position information more accurately. In addition, the spatial resolution of PET is closely related to the anode pixel size. The anode pixel size of the CdZnTe semiconductor detector used in the present invention can reach hundreds of microns, which greatly improves the spatial resolution of the PET system.

[0047] (3) Compared with other detectors, cadmium zinc telluride semiconductor detectors do not require cooling devices or photoelectric conversion devices, which reduces the size of the probe, making the placement of the detector more flexible and easier to place close to the lesion, increasing the detection efficiency of the detector. At the same time, it reduces the power consumption of the instrument and significantly reduces the cost of the PET instrument.

[0048] (4) The three-dimensional position calculation algorithm of gamma rays in the CZT-PET imaging system provided by the present invention can truly realize the three-dimensional imaging of PET technology, and the three-dimensional PET detector technology can greatly improve the spatial resolution of the system.

[0049] (5) The time correction method of the CdZnTe detector provided by the present invention can effectively solve the problem of poor time performance of the CdZnTe detector and improve the time response of the PET system.

[0050] (6) The charge sharing-based interpolation algorithm provided by the present invention compensates for the resolution loss caused by the charge sharing effect, improves the counting rate of the PET system, and also improves the spatial resolution of the PET system. BRIEF DESCRIPTION OF THE DRAWINGS

[0051] Figure 1 It is a schematic diagram of the orthogonal strip and pixel type electrode structure of the cadmium zinc telluride PET detector of the present invention.

[0052] Figure 2 It is a schematic diagram of two cadmium zinc telluride PET modules of the present invention.

[0053] Figure 3 It is a schematic diagram of two single-ring cadmium zinc telluride PET systems of the present invention.

[0054] Figure 4 It is a schematic diagram of two multi-ring cadmium zinc telluride PET systems of the present invention.

[0055] Figure 5 It is a schematic diagram of relevant application scenarios of different types of cadmium zinc telluride PET of the present invention.

[0056] Figure 6 It is a circuit diagram of the filtering and amplifying module of the present invention.

[0057] Figure 7It is a schematic diagram of the method for converting a single CZT-PET detector signal into four-channel signal readout according to the present invention.

[0058] Figure 8 It is a block diagram of the CZT-PET complete system module of the present invention.

[0059] Fig. 9 It is a block diagram of the time correction algorithm of the present invention.

[0060] Fig.10 It is a schematic diagram of the charge sharing interpolation algorithm of the present invention.

[0061] In the figure: 1-ring detector unit; 2-cadmium zinc telluride crystal module; 3-cadmium zinc telluride crystal, 4-strip electrode; 5-square electrode; 6-electrode guard ring. DETAILED DESCRIPTION

[0062] In order to make the purpose, technical solution and advantages of the present invention more clearly understood, the present invention is further described in detail below in conjunction with the accompanying drawings and embodiments. It should be understood that the specific embodiments described herein are only used to explain the present invention and are not intended to limit the present invention.

[0063] Example 1

[0064] The present invention provides a cadmium zinc telluride (CZT) positron emission tomography (PET) imaging system for imaging small animals or human organs, with a cadmium zinc telluride (CZT) semiconductor as the detector body, and each cadmium zinc telluride detection module has a complete preamplifier module, that is, each cadmium zinc telluride detection module includes a cadmium zinc telluride crystal module 2, a charge sensitive preamplifier module and a filter amplifier module. The energy resolution of the cadmium zinc telluride semiconductor detector is excellent, close to 2% for annihilation gamma rays of 511keV, which can greatly improve the accuracy of the ray energy selection range, is conducive to rejecting multiple scattering events, and selecting full energy peak events, so as to achieve the purpose of improving the spatial resolution and signal-to-noise ratio of the system.

[0065] The probe part of the CdZnTe detection module is composed of a whole CdZnTe crystal module 2, which includes a CdZnTe crystal 3 and an electrode. Figure 1 The density of CdZnTe crystal 3 is about 5.8g / cm 3 , with a size of 20mm×5mm×20mm. The crystal electrode type adopts an orthogonal strip electrode structure ( Figure 1 a) or pixel type electrode structure ( Figure 1b), the electrode adopts the evaporation process, which is conducive to making the pixel reach the hundred-micron level and improving the spatial resolution of the PET system; in the orthogonal strip electrode structure, 16 parallel strip electrodes 4 are respectively evaporated on the two opposite sides of the cadmium zinc telluride crystal 3, and the electrode material is metal gold, with a thickness of about 100 microns and a size of 16.4mm×0.9mm. The spacing between adjacent strip electrodes 4 on the same side is 0.1mm. In the pixel type electrode structure, 16×16 square electrodes 5 are evaporated on the anode of one side of the cadmium zinc telluride crystal 3, and the cathode of the other side adopts a planar electrode, and the electrode material is metal gold, with a thickness of about 100 microns, and the size can be selected from 0.9mm, 0.7mm, 0.5mm, 0.3mm, 0.2mm and other sizes, and the spacing between adjacent square electrodes 5 is 0.1mm. In both electrolytic structures, an electrode protection ring 6 is set on the periphery of the electrode to effectively reduce the end surface leakage current of the detector. Circuit boards are soldered on the corresponding sides of the anode and cathode to provide bias and lead-out signals for the semiconductor. The anode is grounded and the cathode is connected to negative high voltage.

[0066] A single CdZnTe PET detector module can be constructed in two ways: a CZT-PET probe module consisting of four CdZnTe crystals laid flat ( Figure 2 a) or a CZT-PET probe module consisting of four stacked CdZnTe crystals ( Figure 2 b) The output signals of four CdZnTe crystals are compressed twice into four signals as the output signals of the CZT-PET single module.

[0067] Multiple CdZnTe detection modules form a ring detector unit 1, forming a single ring PET ( Figure 3 ) or polycyclic PET( Figure 4 ), single-ring PET is capable of full-range imaging of most experimental small animals; in multi-ring PET, multiple ring-shaped detector units are arranged in sequence along the axial direction, and the axial field of view is increased, so full-range imaging can be achieved without moving the target object. Figure 3 A single ring PET is based on a single detector module. Figure 3 The single-ring PET of b is based on 4 detector modules. Both single-ring PET layout structures use 64 detector modules, with a radial field of view of 12 cm, which can perform full-range imaging of most experimental small animals. Multiple ring detector units are arranged in sequence along the axial direction to form a multi-ring PET. Figure 3 a and Figure 3 b Add multiple annular detector units along the axial direction to form a multi-ring PET. Figure 4 a and Figure 4 As shown in b, the axial field of view is increased, and full-range imaging is achieved without moving the target object.

[0068] The radial and axial fields of view of the overall CdZnTe PET can be flexibly increased or decreased according to the size of the target. For small animals, CdZnTe PET can use fewer detector modules, and its radial diameter can be set to about 10 cm ( Figure 5 As shown in the figure above, it can reduce the overall system cost and improve the resolution. For human brain imaging, detectors can be flexibly added according to the shape of the object to be detected, and detection can be performed at different positions. The radial diameter can be set to about 30cm ( Figure 5 For whole-body imaging, the number of detector rings can be increased in the axial direction to improve the sensitivity of the system. The radial diameter can be set to about 70 cm ( Figure 5 as shown in the figure below).

[0069] The output signals of the cadmium zinc telluride crystal 3 are all passed through the charge-sensitive preamplifier. The CZT semiconductor detector realizes ray detection by collecting electron-hole pairs, so a charge-sensitive preamplifier must be selected for signal processing in the subsequent process. Since the output signal of the CZT semiconductor detector is very weak, the charge-sensitive preamplifier used must have superior performance such as high gain and low noise. The charge-sensitive preamplifier used in the present invention has high sensitivity, high signal-to-noise ratio and low power consumption (<150mW), and the pin-type interface facilitates the integration of subsequent circuits. Its operating voltage is ±12V.

[0070] After the CZT pixel signal passes through the preamplifier circuit, the signal amplitude is still only tens of millivolts. The small amplitude electrical signal is easy to cause signal accumulation, resulting in distortion of the collected signal, causing a large number of useful nuclear signals to be lost and distorted, and its baseline is extremely unstable. A large baseline value will cause damage to the back-end data acquisition system. Therefore, the signal processed by the charge sensitive preamplifier needs to be connected to the filter amplifier circuit. The circuit diagram of the filter amplifier module is shown in the figure below. Figure 6 As shown, the filter amplifier circuit adopts a two-stage amplifier circuit to amplify the output signal of the preamplifier, retain the shorter leading edge time of the signal, and greatly reduce the longer falling edge time of the signal, as well as the baseline recovery function, to ensure better matching with the back-end data acquisition system. The filter amplifier module outputs the signal to the signal readout module for further processing.

[0071] A single bar CdZnTe detector outputs 16 anodes and 16 cathodes, a total of 32 signals, and a single pixel type outputs 256 anodes and 1 cathode, a total of 257 signals. Figure 7The signal conversion method shown in the figure converts the output signal of a single CdZnTe detector into four-way signal readout. For the orthogonal strip electrode structure, the X signal and Y signal of the CdZnTe crystal are each converted into two-way signal output; for the pixel electrode structure, all pixel anode signals of the CdZnTe crystal are converted into four-way signal output. Obviously, ordinary data acquisition systems cannot complete the acquisition of so many channels. Signal processing components such as preamplifiers for hundreds or even thousands of channels can be integrated on the ASIC chip. For the axial field of view, detector rings can be added, and ASIC chips can be used for preliminary signal processing. It is necessary to add logic between the detector rings, and then the back end is connected to the ordinary data acquisition system to complete data acquisition. The block diagram of the CZT-PET system module is shown in the figure. Figure 8 shown.

[0072] Example 2

[0073] In order to effectively solve the problems of poor time performance of CdZnTe PET detectors and resolution loss caused by charge sharing effect, the present invention further provides a detector output signal correction algorithm for the CZT-PET imaging system, which includes three aspects: a method for calculating the three-dimensional position of gamma rays, a method for correcting the time resolution, and an interpolation algorithm for charge sharing events.

[0074] 1. Calculation method of the three-dimensional position of gamma rays in CdZnTe PET detectors

[0075] 1. The method for calculating the three-dimensional position of gamma rays of the CdZnTe PET detector (CZT-C32) with orthogonal strip electrode structure is:

[0076] The three-dimensional position (x, y, z) of an interaction event in a CdZnTe crystal is calculated as follows:

[0077]

[0078]

[0079]

[0080] Where G represents the distance between the edge electrode and the edge of the CdZnTe crystal, n represents the maximum induced current at the nth electrode, P represents the width of the anode strip, and A n Indicates the signal peak value at the nth anode, C n represents the signal peak value at the nth cathode, and D represents the thickness of the CdZnTe crystal.

[0081] 2. The method for calculating the three-dimensional position of gamma rays of the CdZnTe PET detector (CZT-P256) with a pixel-type electrode structure is:

[0082] The three-dimensional position (x, y, z) of an interaction event in a CdZnTe crystal is calculated as follows:

[0083]

[0084]

[0085]

[0086] Where G represents the distance between the edge electrode and the edge of the CdZnTe crystal, i and j represent the maximum induced current on the pixel electrode at the i-th row and j-th column, P represents the width of the anode strip, and A ij represents the signal peak value at the anode in the i-th row and j-th column, C represents the signal peak value at the cathode, and D represents the thickness of the CdZnTe crystal.

[0087] 2. Correction Method for the Time Resolution of CdZnTe PET Detector

[0088] Since the time performance of CZT detector is low, the rising edge of cathode signal is long and the rising edge of anode signal is short. For anode signal, the same energy is deposited at different positions, and the signal starting time is different due to the different drift time of electrons. Therefore, the traditional detector output signal time discrimination method will cause large errors. For full energy peak events, the electron drift time and cathode anode ratio maintain a linear relationship at most depths. Then, subtracting the electron drift time from the measured time of the anode signal can obtain a more accurate gamma response time ΔT in CdZnTe detectors. The expression relationship is as follows:

[0089] ΔT=T CZT -T drift , (3-1)

[0090] Among them, T CZT is the readout time of the gamma-ray photon in the CdZnTe detector, T drift is the electron drift time; this correction method can greatly improve the time resolution of the anode signal in the CdZnTe detector and improve the time response of the PET system. The time signal correction block diagram is shown in Fig. 9 shown.

[0091] 3. Interpolation Algorithm for Charge Sharing Events in CdZnTe PET Detectors

[0092] For orthogonal strip and pixel-type anode electrodes of tiny size, when charges approach the anode, induced charges will be generated on multiple anodes, triggering a multi-anode event, which is called charge sharing effect. In PET image reconstruction, the charge sharing event is equivalent to increasing the size of the anode pixel, thus reducing the spatial resolution of the PET system. In order to solve this problem, the present invention proposes the following interpolation algorithm:

[0093] The charge cloud is detected by multiple pixels. According to the relationship that the signal amplitude of adjacent anode pixels is proportional to the charge collected, the signal amplitudes of three adjacent and sequentially arranged anode pixels are A and B, respectively. 1 , A 2 , A 3 , calculate the ratio R of the signal amplitudes of two adjacent anode electrodes:

[0094] R 21 =A 2 / A 1 ; (4-1)

[0095] R 23 =A 2 / A 3 ; (4-2)

[0096] According to the ratio R 21 and R 23 , A 2 The corresponding anode pixel range is divided into five regions, respectively recorded as regions A, B, C, D, and E, and the division basis is as follows:

[0097] 1<R 21 <3, area A;

[0098] 3<R 21 <9, area B;

[0099] 9<R 21 or 9<R 23 , is region C;

[0100] 3<R 23 <9, it is area D;

[0101] 1<R 23 <3, it is area E;

[0102] like Fig.10 As shown in (a), the size of each divided area is one-fifth of the sum of the anode pixel and the anode gap, which is equivalent to using the interpolation algorithm of shared events to further reduce the anode pixel size to one-fifth, thereby determining the three-dimensional position of the charge sharing event and improving the position resolution of the charge sharing event. At the same time, the counting rate of each area is basically consistent, thereby also improving the counting rate of the system. The division of each area is based on basically ensuring that the number of hit events in the area is consistent with the number of hit events in the adjacent area. The division principle is as follows Fig.10 (b) as shown.

[0103] The above description is only a preferred embodiment of the present invention and is not intended to limit the present invention. Any modifications, equivalent substitutions and improvements made within the spirit and principles of the present invention should be included in the protection scope of the present invention.

Claims

1. A method for calculating the three-dimensional position of a detector gamma ray of a cadmium zinc telluride positron emission tomography system, characterized in that: The CdZnTe positron emission tomography system uses a CdZnTe detection module to form a ring detector unit, forming a single-ring PET or a multi-ring PET. The CdZnTe detection module includes a CdZnTe crystal module, a charge-sensitive preamplifier module and a filter amplifier module. The CdZnTe crystal module includes a CdZnTe crystal and an electrode. The electrode type of the CdZnTe crystal adopts an orthogonal strip electrode structure or a pixel-type electrode structure. The output signal of the CdZnTe crystal is connected to the charge-sensitive preamplifier module through a circuit, the charge-sensitive preamplifier module is connected to the filter amplifier module through a circuit, and the filter amplifier module is connected to the signal readout module through a circuit. The calculation method is as follows: (1) The method for calculating the three-dimensional position of gamma rays in a CdZnTe PET detector with an orthogonal strip electrode structure is: For the three-dimensional location of the interaction events occurring in the CdZnTe crystal ( x , y , z ) is calculated as follows: ，（1-1） ，（1-2） ，（1-3） in, G represents the distance between the edge electrode and the edge of the CdZnTe crystal, n Indicates that the trigger n The induced current is the largest at the electrode. P represents the width of the anode strip, A n Indicated in n The signal peak on the anode is C n represents the signal peak at the nth cathode, D Indicates the thickness of CdZnTe crystal; (2) The method for calculating the three-dimensional position of gamma rays in a CdZnTe PET detector with a pixel-type electrode structure is: For the three-dimensional location of the interaction events occurring in the CdZnTe crystal ( x , y , z ) is calculated as follows: ，（2-1） ，（2-2） ，（2-3） in, G represents the distance between the edge electrode and the edge of the CdZnTe crystal, i , j Respectively indicate the trigger i Line j The induced current on the pixel electrode of the column is the largest. P represents the width of the anode strip, A ij Indicated in i Line j The signal peak on the column anode, C represents the signal peak at the cathode, D Indicates the thickness of the CdZnTe crystal.

2. The calculation method according to claim 1, characterized in that: In the orthogonal strip electrode structure, 16 parallel strip electrodes are respectively deposited on two opposite sides of the CdZnTe crystal; in the pixel type electrode structure, 16×16 square electrodes are deposited on the anode on one side of the CdZnTe crystal, and a planar electrode is used as the cathode on the other side.

3. The calculation method according to claim 2, characterized in that: In the CdZnTe crystal module, several charge-sensitive preamplifiers are integrated through an ASIC chip.

4. The calculation method according to claim 2, characterized in that: The electrodes are made of gold material, and the spacing between adjacent strip electrodes on the same side in the orthogonal strip electrode structure and the spacing between adjacent square electrodes in the pixel type electrode structure are both 0.1 mm.

5. The calculation method according to claim 1, characterized in that: An electrode protection ring is arranged around the periphery of the electrode.

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