阅读说明：本技术 放射线检测装置以及具备该放射线检测装置的核医学诊断装置 (Radiation detection device and nuclear medicine diagnosis device provided with same ) 是由 中泽诚之 古宫哲夫 于 2017-07-03 设计创作，主要内容包括：放射线检测装置(300)被用于核医学诊断装置,具备多个闪烁体(44)、半导体受光器件(SiPM)、位置检测电路(214)以及定时检测电路(216)。闪烁体将从被检体(15)照射的γ射线变换为荧光。半导体受光器件与多个闪烁体分别对应地设置,将由对应的闪烁体进行变换所得到的荧光变换为电信号。位置检测电路基于来自半导体受光器件的电信号来确定多个闪烁体中的γ射线的检测位置。定时检测电路与半导体受光器件的阳极连接,用于确定与检测γ射线的事件的发生时间对应的时间信息。(A radiation detection device (300) is used in a nuclear medicine diagnosis device, and is provided with a plurality of scintillators (44), semiconductor light receiving devices (SiPM), a position detection circuit (214), and a timing detection circuit (216). The scintillator converts gamma rays irradiated from the subject (15) into fluorescence. The semiconductor light receiving device is provided in correspondence with each of the plurality of scintillators, and converts fluorescence converted by the corresponding scintillator into an electric signal. The position detection circuit determines the detection position of the gamma ray in the plurality of scintillators based on the electric signal from the semiconductor light receiving device. The timing detection circuit is connected to the anode of the semiconductor light receiving device and determines time information corresponding to the occurrence time of an event for detecting gamma rays.)

1. A radiation detection apparatus used in a nuclear medicine diagnosis apparatus, the radiation detection apparatus comprising:

a plurality of scintillators that convert gamma rays irradiated from a subject into fluorescence;

semiconductor light receiving devices provided in correspondence with the plurality of scintillators, respectively, and converting fluorescence converted by the corresponding scintillators into electric signals;

a position detection circuit that determines detection positions of gamma rays in the plurality of scintillators based on an electric signal from an anode of the semiconductor light receiving device; and

and a timing detection circuit connected to the anode of the semiconductor light receiving device, for determining time information corresponding to the occurrence time of an event for detecting gamma rays.

2. The radiation detecting apparatus according to claim 1,

the plurality of scintillators are arranged in an array,

the position detection circuit determines a detection position of a gamma ray in the array based on a first weighted addition signal of electric signals with respect to a row of the array and a second weighted addition signal of electric signals with respect to a column of the array.

3. The radiation detection apparatus according to claim 2, further comprising:

a supply voltage; and

a center of gravity operation circuit that generates the first weighted addition signal and the second weighted addition signal,

a plurality of the semiconductor light receiving devices are connected in parallel between the power supply voltage and the center-of-gravity operation circuit,

the cathode of the semiconductor light receiving device is connected with the power supply voltage, the anode is connected with the gravity center operation circuit,

the timing detection circuit determines the time information when a signal from any one of the plurality of semiconductor light receiving devices is detected.

4. The radiation detecting apparatus according to claim 3,

the timing detection circuit further includes a capacitor connected between the timing detection circuit and the anode of each of the semiconductor light receiving devices.

5. The radiation detecting apparatus according to claim 4,

the capacitance of the capacitor is determined by the number of the semiconductor light receiving devices connected in parallel to the center-of-gravity operation circuit.

6. The radiation detecting apparatus according to claim 3,

the number of the semiconductor light receiving devices connected in parallel to the center-of-gravity operation circuit is determined in accordance with the detection accuracy required by the timing detection circuit.

7. A nuclear medicine diagnostic apparatus provided with the radiation detection apparatus according to any one of claims 1 to 6.

Technical Field

The present invention relates to a radiation detection apparatus for detecting radiation emitted from a subject to which a radiopharmaceutical is administered, and a nuclear medicine diagnostic apparatus including the radiation detection apparatus, and more particularly, to a technique for improving the detection time and detection position accuracy of gamma rays detected by the radiation detection apparatus.

Background

Nuclear medicine imaging devices such as a PET (Positron Emission Tomography) device and a SPECT (Single Photon Emission Computed Tomography) device are known as nuclear medicine diagnostic devices that obtain medical data of a subject based on radiation emitted from the subject to which a radiopharmaceutical has been administered.

A PET apparatus detects two gamma rays generated by annihilation of positrons (positrons) using a plurality of detectors. Specifically, a radiopharmaceutical (radiotracer) containing a positron-emitting nuclide is administered to a subject, and annihilation gamma rays emitted from the administered subject are detected by a plurality of radiation detectors. When the two radiation detectors detect the gamma rays for a predetermined time, the two gamma rays are counted as a pair of annihilation gamma rays, and the generation position of the annihilation gamma rays is determined on a straight line connecting the two radiation detection positions at which the gamma rays are detected. In a TOF (Time of Flight) type PET apparatus, a nuclear medicine image can be obtained by determining a generation point of pair annihilation gamma rays on the straight line by using a Time difference between the pair annihilation gamma rays detected by the two detectors and imaging a dose distribution of the detected gamma rays.

In a nuclear medicine diagnostic apparatus such as a PET apparatus, a scintillator that converts incident gamma rays into fluorescence having a peak in an ultraviolet region and a light receiving device that converts photoelectrons from the scintillator into an electric signal after multiplying the photoelectrons are generally used as a detector for detecting gamma rays. As the light receiving device, a Photomultiplier Tube (PMT) using a plurality of dynodes has been conventionally used, but in recent years, a silicon Photomultiplier Tube (SiPM) using an Avalanche Photodiode (APD) as a semiconductor element array has been used in some cases. Since sipms have characteristics less susceptible to magnetic influences than PMTs, sipms are also applicable to devices in which an MRI apparatus and a nuclear medicine diagnostic apparatus are integrated.

In this specification, a device that converts gamma rays into an electrical signal is referred to as a "radiation detector (or gamma ray detector)", and a configuration including the radiation detector and a signal processing circuit at a subsequent stage is referred to as a "radiation detection apparatus".

Disclosure of Invention

Problems to be solved by the invention

In a nuclear medical diagnostic apparatus, it is required to provide an image with higher resolution. In order to meet this demand, a radiation detector having a plurality of light receiving devices arranged in an array may be used. With such a configuration, the incident position of the gamma ray in the detectable region of the radiation detector can be detected with higher accuracy.

In such a configuration, in order to improve the position detection accuracy of the radiation detector, it is more preferable to provide a separate readout circuit for each of the plurality of light receiving devices. However, when a plurality of light receiving devices are arranged in a two-dimensional array in one radiation detector (for example, 100 × 100), several thousands to several tens of thousands of read-out circuits are required for each detector, and several times to several tens of times of read-out circuits are required for the entire PET apparatus. Thus, there is a fear that the scale of the apparatus and the cost of the apparatus increase greatly.

In order to solve such a problem, non-patent document 1 proposes the following method: the input position of a gamma ray in a detectable region is determined using a small number of readout circuits by connecting output signals of a plurality of SiPMs in parallel to one readout circuit using a multiplexer circuit and performing a gravity center calculation of each output signal.

Here, when the SiPM is used as the light receiving device, the low-pass filter is formed by the parasitic capacitance of the SiPM itself and the input impedance of the signal processing circuit, and therefore, the high-frequency component of the received light signal may be deteriorated. In particular, when the number of sipms connected in parallel is increased, the degree of deterioration of the high-frequency component of the light reception signal is further increased, and therefore it is difficult to detect the rising edge of the light reception signal. Therefore, detection characteristics of the incidence timing of the γ -ray (hereinafter, also referred to as "timing characteristics") may be degraded.

The present invention has been made to solve the above-described problems, and an object of the present invention is to improve position detection accuracy while suppressing a decrease in timing characteristics of detection signals in a radiation detection apparatus used in a nuclear medicine diagnostic apparatus.

Means for solving the problems

The radiation detection apparatus of the present invention is used for a nuclear medicine diagnosis apparatus. The radiation detection apparatus includes a plurality of scintillators, a semiconductor light receiving device, a position detection circuit, and a timing detection circuit. The scintillator converts gamma rays irradiated from the subject into fluorescence. The semiconductor light receiving device is provided in correspondence with each of the plurality of scintillators, and converts fluorescence obtained by conversion by the corresponding scintillator into an electric signal. The position detection circuit determines the detection positions of the gamma rays in the plurality of scintillators based on an electric signal from the anode of the semiconductor light receiving device. The timing detection circuit is connected to the anode of the semiconductor light receiving device and determines time information corresponding to the occurrence time of an event for detecting gamma rays.

In this way, the same electrical signal from the anode of the semiconductor light receiving device can be used to specify the light receiving position of the gamma ray and the detection timing of the gamma ray. Therefore, by adjusting the electric signal from the anode, the characteristics of both the position detection and the time detection can be changed in the same tendency. Therefore, the improvement of the position detection accuracy and the improvement of the time detection accuracy can be achieved at the same time.

Preferably, the plurality of scintillators are arranged in an array. The position detection circuit determines a detection position of a gamma ray in the array based on a first weighted addition signal of the electric signals with respect to the rows of the array and a second weighted addition signal of the electric signals with respect to the columns of the array.

In this way, the detection position of the gamma ray can be specified based on the signals obtained by summing up the plurality of scintillators arranged in an array for each row and each column. Therefore, in the position detection circuit, the number of signals used for position detection can be smaller than the number of scintillators (i.e., the number of semiconductor light receiving devices), and an increase in the scale and cost of the apparatus can be suppressed.

Preferably, the radiation detection apparatus further includes a power supply voltage and a center-of-gravity operation circuit that generates the first weighted addition signal and the second weighted addition signal. The plurality of semiconductor light receiving devices are connected in parallel between a power supply voltage and the center-of-gravity operation circuit. The cathode of the semiconductor light receiving device is connected to a power supply voltage, and the anode is connected to a center-of-gravity arithmetic circuit. When a signal from any one of the plurality of semiconductor light receiving devices is detected, the timing detection circuit specifies time information corresponding to the occurrence time of an event for detecting gamma rays.

With this configuration, when the gamma ray is irradiated to any one of the plurality of semiconductor light receiving devices connected in parallel, the gamma ray can be appropriately detected by the timing detection circuit.

Preferably, the timing detection circuit further includes a capacitor connected between the timing detection circuit and the anode of each semiconductor light receiving device.

By connecting the timing detection circuit to the anode of the semiconductor light receiving device via the capacitor in this manner, a high-frequency component (i.e., a component having a high response speed) of the electric signal of the semiconductor light receiving device can be detected by the timing detection circuit. Therefore, the time detection accuracy of the γ -ray can be improved.

Preferably, the capacitance of the capacitor is determined according to the number of semiconductor light receiving devices connected in parallel to the center-of-gravity operation circuit.

When the semiconductor light receiving devices are connected in parallel, a low-pass filter is formed by a parasitic capacitance component thereof and a capacitor connected to the timing detection circuit. The low-pass filter formed attenuates the high-frequency component of the electric signal detected by the timing detection circuit, which may result in a decrease in time detection accuracy. Therefore, by appropriately setting the capacitance of the capacitor connected to the timing detection circuit in accordance with the number of semiconductor light receiving devices connected in parallel, it is possible to suppress a decrease in time detection accuracy.

Preferably, the number of semiconductor light receiving devices connected in parallel to the center-of-gravity operation circuit is determined in accordance with detection accuracy required by the timing detection circuit.

As described above, when the semiconductor light receiving devices are connected in parallel, the time detection accuracy of the timing detection circuit is affected by the low-pass filter formed by the parasitic capacitance component thereof. Therefore, the desired time detection accuracy can be ensured by determining the number of semiconductor light receiving devices connected in parallel according to the detection accuracy required by the timing detection circuit.

A nuclear medicine diagnosis device of the present invention includes the radiation detection device described in any one of the above.

ADVANTAGEOUS EFFECTS OF INVENTION

According to the present invention, in a radiation detection apparatus used for a nuclear medicine diagnosis apparatus, it is possible to suppress a decrease in timing characteristics of detection signals and to improve position detection accuracy.

Drawings

Fig. 1 is an overall schematic diagram of a PET apparatus according to the present embodiment.

Fig. 2 is a schematic perspective view of the gamma-ray detector in fig. 1.

Fig. 3 is a functional block diagram showing details of the data collection unit in fig. 1.

Fig. 4 is a diagram for explaining the waveform shaping circuit of fig. 3.

Fig. 5 is a diagram for explaining a signal processing circuit in a comparative example.

Fig. 6 is a diagram for explaining an example of the configuration of the center-of-gravity calculation circuit.

Fig. 7 is a diagram showing an example of a position detection map detected in the comparative example.

Fig. 8 is a diagram showing an ideal position detection map.

Fig. 9 is a diagram showing an example of a centroid calculation waveform in the comparative example.

Fig. 10 is a diagram for explaining a signal processing circuit in the present embodiment.

Fig. 11 is a diagram showing an example of a centroid calculation waveform in the present embodiment.

Fig. 12 is a diagram showing an example of a position detection map in the present embodiment.

Fig. 13 is a diagram showing another example of a position detection map in the present embodiment.

Detailed Description

Embodiments of the present invention will be described in detail below with reference to the drawings. In the drawings, the same or corresponding portions are denoted by the same reference numerals, and description thereof will not be repeated.

[ Structure of Nuclear medicine diagnostic apparatus ]

Fig. 1 is an overall schematic diagram of a nuclear medical diagnostic apparatus according to the present embodiment. Fig. 1 shows an example in which the nuclear medicine diagnosis apparatus is the PET apparatus 100, but the nuclear medicine diagnosis apparatus is not limited to this, and may be any other apparatus as long as it uses a so-called radiation detector, for example, a SPECT apparatus. In the present embodiment, a case of gamma rays will be described as an example of radiation.

Referring to fig. 1, the PET apparatus 100 includes a gantry apparatus 10, a control apparatus 200, a display apparatus 260, and an operation unit 270. Fig. 1 (a) is a front view of the stage device 10, and (b) is a side view of the stage device 10.

The gantry apparatus 10 includes a top plate 20 on which the subject 15 is placed, a moving device 22 that moves the top plate 20, a substantially cylindrical gantry 30 having an opening, and a detector ring 40 disposed in the gantry 30.

The control device 200 includes a data collection unit 210, a control unit 220, and a drive unit 230. The control unit 220 includes a storage device such as a CPU (central processing unit) or a memory. The data collection unit 210 and the drive unit 230 may be configured by a microprocessor or an FPGA (Field Programmable Gate Array), or may be a part of the CPU of the control unit 220.

The subject 15 is placed on a cushion pad 24 provided on the top plate 20. The top plate 20 is provided so as to penetrate through the opening portions of the gantry 30 and the detector ring 40 in the Z direction indicated by an arrow AR in the drawing, and is capable of reciprocating in the Z direction. The movement device 22 is controlled by a drive signal from the drive unit 230, and the movement device 22 adjusts the height of the top plate 20 and moves the top plate 20 in the Z direction, thereby introducing the subject 15 placed on the top plate 20 into the opening of the gantry 30.

The detector ring 40 is configured by arranging a plurality of unit rings in the Z direction, the unit rings being formed by radially arranging a plurality of radiation detectors 42 on a plane perpendicular to the Z direction.

As shown in fig. 2, the radiation detector 42 (hereinafter also referred to as "γ -ray detector 42") includes a light receiving sensor 45 and a scintillator block 44 including a plurality of scintillators in an array. Each scintillator of the scintillator block 44 converts radiation (γ -rays) 52 emitted from a radiopharmaceutical (radiotracer) 50 (for example, Fluorodeoxyglucose (FDG)) containing a positron-emitting nuclide administered to the subject 15 into fluorescence having a peak in the ultraviolet region. The light receiving sensor 45 is provided with light receiving devices corresponding to the scintillators, and each light receiving device multiplies photoelectrons converted by the corresponding scintillator and converts the multiplied photoelectrons into an electric signal. In this embodiment, as the light receiving device, a silicon photomultiplier (SiPM) using an avalanche photodiode (AMD) as a semiconductor array is used. The gamma-ray detector 42 outputs the generated electric signal to the data collection unit 210 in the control device 200 of fig. 1.

The data collection unit 210 processes the signal received from the gamma-ray detector 42 and outputs the processed signal to the control unit 220. The control unit 220 images the dose distribution of the detected gamma rays based on the reception signal from the data collection unit 210, and displays the image on the display device 260.

The operation unit 270 includes a pointing device such as a keyboard, a touch panel, or a mouse (none of which are shown). The operation unit 270 receives an instruction to operate the moving device 22 of the stage device 10 and an instruction to start/stop imaging from the operator. The operation unit 270 outputs a signal generated in response to an operation by the operator to the control unit 220. The control unit 220 controls the driving unit 230 based on a signal from the operation unit 270, thereby driving the moving device 22.

In the present embodiment, the configuration including the gamma-ray detector 42 and the data collection unit 210 is referred to as a "radiation detection apparatus".

Fig. 3 is a functional block diagram for explaining details of the data collection unit 210 in fig. 1. Referring to fig. 3, the data collection unit 210 includes a Front End (FE) circuit 212 and a coincidence counting circuit 218 provided corresponding to each of the gamma-ray detectors 42 constituting the detector ring 40. Each FE circuit 212 includes a waveform shaping circuit 213, a position detection circuit 214, an energy detection circuit 215, and a timing detection circuit 216.

The waveform shaping circuit 213 receives the electrical signal generated by the gamma-ray detector 42, and performs waveform shaping processing of analog waveform data of the electrical signal. Specifically, the waveform shaping circuit 213 performs arithmetic processing such as integration processing and differentiation processing on the analog waveform data from the γ -ray detector 42 as shown in fig. 4 (a), and generates data having peak values representing energy as shown in fig. 4 (b). The waveform shaping circuit 213 outputs the generated data to the position detection circuit 214 and the energy detection circuit 215.

The position detection circuit 214 receives the data generated by the waveform shaping circuit 213, and determines in which scintillator in the scintillator block 44 the gamma ray is detected. Specifically, the position of the scintillator in which the γ -ray is detected is determined by calculating the position of the center of gravity of the data generated by the waveform shaping circuit 213. The position detection circuit 214 outputs data indicating the determined scintillator position to the coincidence counting circuit 218.

The energy detection circuit 215 receives the data generated by the waveform shaping circuit 213 and detects energy. The energy detection circuit 215 outputs data representing the detected energy to the coincidence counting circuit 218.

The timing detection circuit 216 detects time information corresponding to the occurrence time of an event for detecting a gamma ray, that is, the detection time (incidence timing) of a gamma ray, from the analog waveform data from the gamma ray detector 42 shown in fig. 4 (a). For example, a time point at which the value of the analog waveform data shown in fig. 4 (a) exceeds a predetermined threshold value is determined as the detection time of the γ -ray. The timing detection circuit 216 outputs the data of the determined detection time to the coincidence counting circuit 218.

The coincidence counting circuit 218 receives data from each FE circuit 212, and generates coincidence counting information for determining the incidence direction of pair annihilation gamma rays emitted from positrons. Specifically, the coincidence counting circuit 218 searches for a combination of the gamma ray detectors in which the incident timing (detection time) of the gamma ray is within a predetermined time window width and the energy of the light reception signal is within a fixed energy window width based on the data from the plurality of FE circuits 212. Then, the coincidence counting circuit 218 determines a combination of the retrieved γ -ray detectors as the γ -ray detectors that simultaneously detect two annihilation photons emitted from one positron. That is, it means that the radioactive tracer 50 emitting the gamma ray exists on the straight line connecting the two determined gamma ray detectors.

The coincidence counting circuit 218 calculates the time difference between the two annihilation photons from the radiotracer 50 (i.e., the distance from the gamma-ray detector to the radiotracer: TOF) based on the data from the timing detection circuit 216, and determines the position of the radiotracer 50 on the straight line connecting the two gamma-ray detectors. The coincidence counting circuit 218 outputs data on the determined position of the gamma ray detector and the position of the radioactive tracer 50 to the control unit 220.

The control unit 220 generates an image of the subject 15 by reconstructing the data received from the coincidence counting circuit 218, and displays the image on the display device 260. A diagnostician such as a doctor performs nuclear medicine diagnosis using the displayed image of the subject 15.

In the PET apparatus having such a configuration, it is required to provide an image with a higher resolution for accurate diagnosis. In order to improve the resolution, it is necessary to (1) improve the incident position detection accuracy of the gamma rays in each gamma ray detector and (2) improve the detection accuracy (time resolution) of the time difference of detection of the gamma rays in the two gamma ray detectors.

In order to improve the detection accuracy of the incident position in each gamma-ray detector, it is preferable to provide a separate signal processing circuit for each light receiving device (SiPM) at a ratio of 1: 1. However, when a plurality of light receiving devices are arranged in a two-dimensional array in one γ -ray detector, the same number of signal processing circuits as the number of light receiving devices are required, and as a result, several tens of thousands to several hundreds of thousands of signal processing circuits may be required in the entire PET apparatus. Thus, the scale of the apparatus becomes enormous, and the apparatus cost also increases.

To cope with such a situation, the following configuration is proposed: in each gamma-ray detector, a plurality of light receiving devices are connected in parallel using a multiplexer circuit, and one signal processing circuit is provided for the plurality of light receiving devices, thereby reducing the number of signal processing circuits in the entire apparatus.

Fig. 5 is a diagram for explaining a signal processing circuit in the radiation detection apparatus 300A of the comparative example. Referring to fig. 5, the signal processing circuit of this comparative example includes a BIAS voltage BIAS, a resistance RL, a plurality of sipms as light receiving devices, a center-of-gravity operation circuit 60, and a capacitor Cf.

The plurality of sipms are provided corresponding to the plurality of scintillators (fig. 2) arranged in an array, and are arranged in an array as shown in fig. 6. In the example of fig. 6, a case where 16 sipms are configured as a 4 × 4 array is shown.

One end of the resistor RL is connected to the BIAS, and the other end is connected to the cathode of each SiPM of the plurality of sipms. The anodes of the sipms are connected to the center-of-gravity calculation circuit 60.

One end of the resistor RL (the cathode of SiPM) is also connected to the timing detection circuit 216 of the FE circuit 212 via the capacitor Cf. Only the high-frequency component of the voltage change at the anode of the SiPM (i.e., the node ND) is transmitted to the timing detection circuit 216 through the capacitor Cf.

When a gamma ray is incident on the gamma ray detector 42 and is detected by any one of the plurality of sipms, a current flows through the SiPM that has detected the gamma ray, and the voltage at the node ND decreases in a pulse shape as shown in fig. 2 (a). If the amount of decrease in the voltage is larger than a predetermined amount, a signal is output to the timing detection circuit 216. Thereby, the detection time of the γ -ray is determined in the timing detection circuit 216.

The center of gravity calculation circuit 60 is a circuit that generates a signal for determining which SiPM among a plurality of sipms arranged in an array detects a gamma ray. Specifically, as shown in fig. 6, the parallel signals of each row of sipms arranged in an array are weighted by resistors R1 to R4, and added to generate position detection signals Xa and Xb. Similarly, the parallel signals of the respective columns of the SiPM are weighted by the resistors R1 to R4, respectively, and then added to generate the position detection signals Ya and Yb. The generated position detection signal is output to the waveform shaping circuit 213 of the FE circuit 212.

For example, when the resistance value is set to R1< R2< R3< R4, the amplitude of the position detection signal Xa is maximum when a γ ray is detected in the SiPM in the first row and minimum when a γ ray is detected in the SiPM in the fourth row. That is, it is possible to determine which row of sipms the gamma ray is detected in, based on the amplitude of the position detection signal Xa. Similarly, it is possible to determine in which column of sipms the gamma ray is detected, based on the amplitude of the position detection signal Ya. Thus, by using the position detection signal Xa and the position detection signal Ya, it is possible to determine which of the sipms arranged in an array the γ -ray is detected in (the position detection circuit 214).

In addition, the position detection signals Xb and Yb are resistively connected in the order reverse to the order of resistance connection in the case of the position detection signals Xa and Ya for each row and column. Therefore, the tendency of the amplitude levels of the position detection signals Xb and Yb is opposite to the tendency of the amplitude levels of the position detection signals Xa and Ya, respectively. For example, when a γ ray is detected in the SiPM in the first column, the amplitude of the position detection signal Xb is minimum, and when a γ ray is detected in the SiPM in the fourth row, the amplitude of the position detection signal Xb is maximum. By using such a position detection signal having an opposite tendency, for example, even when the entire signal is shifted, the SiPM in which the gamma ray is detected can be accurately specified.

Here, in the signal processing circuit of the comparative example shown in fig. 5, in order to improve the detection accuracy of the detection time of the γ -ray, it is necessary to increase the voltage variation of the node ND, and therefore, it is necessary to increase the resistance RL connected to the BIAS voltage BIAS. However, in this case, since it takes time to charge the capacitor Cf after the detection of the γ -rays by the SiPM is completed, the response of the position detection signal in the center-of-gravity calculation circuit 60 becomes sluggish. Thus, the position detection accuracy may be degraded.

On the other hand, if the resistance RL is reduced in order to improve the position detection accuracy, the voltage drop amount of the node ND becomes smaller, and therefore the detection time accuracy (time resolution) of the γ -ray is reduced. That is, in the signal processing circuit shown in fig. 5, the position detection accuracy and the time resolution are in a trade-off relationship, and it is difficult to improve the accuracy of both.

Fig. 7 shows an example of a position detection map of gamma rays obtained by using a 4 × 4 SiPM array when signal processing is performed using this comparative example. Although 16 sipms can be distinguished in fig. 7, the map of fig. 7 is a map in which a large deformation occurs as compared with the ideal position detection map shown in fig. 8.

In the comparative example, fig. 9 shows an example of the position detection signals Xa, Xb, Ya, Yb output from the center-of-gravity calculation circuit 60. As can be seen from fig. 9, when the signal processing circuit of the comparative example is used, some of the position detection signals (L2 (Xb) and L4(Yb) in fig. 9) have negative voltage amplitude. Therefore, the detection position in the position detection circuit 214 cannot be correctly determined.

In order to solve the above problem, the present embodiment employs the following configuration: a timing signal for determining the detection time of the gamma ray is read from the anode side of each SiPM without using the resistance RL in the signal processing circuit of fig. 5. With such a configuration, it is possible to prevent a decrease in the responsiveness of the position detection signal and a decrease in the level of the timing signal due to the resistance RL, and thus it is possible to improve both the position detection accuracy and the time resolution.

Fig. 10 is a diagram for explaining a signal processing circuit of the radiation detection apparatus 300 according to the present embodiment. Referring to fig. 10, in the signal processing circuit, the cathodes of the plurality of sipms are not connected to the BIAS voltage BIAS via a resistor, and the anodes of the sipms are individually connected to the center of gravity calculation circuit 60. The anodes of the sipms are connected in parallel via a capacitor Cf, and further connected to the timing detection circuit 216 of the FE circuit 212.

In this way, since there is no resistance component (BIAS resistance) connected to the BIAS voltage BIAS, it is possible to suppress a decrease in the responsiveness of the position detection signal in the center of gravity calculation circuit 60. Further, since the timing signal is acquired from the anode side of the SiPM, a voltage difference between when the gamma ray is detected and when the gamma ray is detected can be sufficiently secured even without the offset resistor, and thus a decrease in the level of the timing signal can be suppressed. This can improve the position detection accuracy and the detection time accuracy (time resolution).

In this case, when the impedance of the timing signal readout circuit (FE circuit 212) is increased by the parasitic capacitance component of SiPM, the high frequency component of the timing signal is cut off by the low pass filter, and there is a possibility that signal degradation occurs. Therefore, a timing signal detection circuit is desired to have a configuration in which the input impedance is as low as possible.

The capacitance of the capacitor Cf is also preferably determined by the parasitic capacitance component of SiPM. In other words, it is desirable to determine the capacitance of the capacitor Cf in accordance with the number of sipms connected in parallel to the center-of-gravity calculation circuit.

Alternatively, the capacitance of the predetermined capacitor Cf may be determined by the number of sipms connected in parallel according to the allowable deterioration state of the timing signal (i.e., the detection accuracy required by the timing detection circuit).

Fig. 11 shows an example of the position detection signals Xa, Xb, Ya, Yb output from the centroid calculating circuit 60 in the case where the radiation detection apparatus according to the present embodiment is used. In fig. 11, unlike fig. 9 of the comparative example, the amplitudes of all the position detection signals Xa, Xb, Ya, Yb are positive voltage values. Therefore, by using the signal processing circuit of the present embodiment, the distortion of the position detection map as shown in fig. 7 in the comparative example can be reduced.

Fig. 12 is a diagram showing an example of a position detection map of gamma rays obtained by a 4 × 4 SiPM array in the case where the radiation detection apparatus according to the present embodiment is used. The position detection map of fig. 12 is a map that is reduced in distortion compared to the position detection map of fig. 7 shown in the comparative example and is closer to the ideal position detection map shown in fig. 8. As can be seen from fig. 12, when the radiation detection apparatus according to the present embodiment is used, the positional relationship of the sipms arranged in the grid shape can be accurately recognized.

Fig. 13 is an example of a position detection map in the case where the number of sipms is further increased, and fig. 13 uses a total of 144 sipms of 12 × 12. In fig. 13, although there are some variations, the positions of sipms arranged in a grid pattern can be accurately grasped as a whole.

As described above, by using the radiation detection apparatus according to the present embodiment, it is possible to improve the position detection accuracy and the time resolution of the radiation emitted from the subject in the nuclear medicine diagnosis apparatus.

The presently disclosed embodiments are to be considered in all respects as illustrative and not restrictive. The scope of the present invention is defined by the claims, rather than the description above, and includes all modifications within the meaning and range equivalent to the claims.

Description of the reference numerals

10: a stage device; 15: a subject; 20: a top plate; 22: a mobile device: 24: a cushion pad; 30: a frame; 40: a detector ring; 42: a gamma ray detector; 44: a scintillator block; 45: a light receiving sensor; 50: a radioactive tracer; 60: a center of gravity operation circuit; 100: a PET device; 200: a control device; 210: a data collection unit; 212: an FE circuit; 213: a waveform shaping circuit; 214: a position detection circuit; 215: an energy detection circuit; 216: a timing detection circuit; 218: a coincidence counting circuit; 220: a control unit; 230: a drive section; 260: a display device; 270: an operation section; 300. 300A: a radiation detection device; BIAS: a bias voltage; cf: a capacitor; ND: and (4) nodes.