JP7582804B2 - Nuclear Medicine Diagnostic Equipment - Google Patents

Description

The embodiments disclosed in this specification and the drawings relate to a nuclear medicine diagnostic device.

Traditionally, PET (Positron Emission Computed Tomography) devices are known as nuclear medicine diagnostic devices capable of performing functional diagnosis of a subject's biological tissue. In PET devices, semiconductor gamma ray detectors are used. For example, semiconductor gamma ray detectors have a scintillator and a photomultiplier. The scintillator converts gamma rays emitted from the internal tissue of the subject into light having a peak in the ultraviolet region and outputs the light. The photomultiplier converts the light output from the scintillator into an electrical signal. Silicon photomultipliers (SiPMs) are used as photomultipliers. For example, in the detector of a PET device, a SiPM is provided for each scintillator.

Here, the time resolution of the SiPM is affected by the magnitude of the reverse bias voltage applied to the SiPM. When the temperature of the SiPM rises, for example, the value of the reverse bias voltage is increased to adjust the rise of the signal. However, increasing the value of the reverse bias voltage is one of the factors that causes the temperature of the SiPM to rise further. In addition, the time resolution may deteriorate exponentially when the temperature of the SiPM reaches a certain level or higher. In fact, it is conceivable that the circuit itself may not be able to keep up with the increase in power consumption.

In addition, when imaging with a PET device, the gamma ray count rate changes depending on the subject's physique, the measurement site, the type of drug administered to the subject, and other factors. Here, it is thought that the optimal value of the reverse bias voltage applied to the SiPM differs depending on the count rate. However, when imaging with a PET device, the value of the reverse bias voltage applied to the SiPM is constant regardless of the gamma ray count rate.

US Patent Application Publication No. 2011/0248175 US Patent Application Publication No. 2012/0305784 US Patent Application Publication No. 2017/0276808

One of the problems that the embodiments disclosed in this specification and the drawings attempt to solve is to achieve both the inherent performance and stability of a semiconductor gamma ray detector. However, the problems that the embodiments disclosed in this specification and the drawings attempt to solve are not limited to the above problem. Problems corresponding to the effects of each configuration shown in the embodiments described below can also be positioned as other problems.

The nuclear medicine diagnosis device according to the embodiment includes a semiconductor gamma ray detector, an acquisition unit, and a control unit. The semiconductor gamma ray detector counts incident gamma rays. The acquisition unit acquires information related to the gamma ray count rate. The control unit controls the bias voltage value applied to the semiconductor gamma ray detector based on the information related to the count rate.

FIG. 1 is a block diagram showing an example of the configuration of a PET apparatus according to an embodiment. FIG. 2 is a diagram illustrating an example of the configuration of a detector module according to the embodiment. FIG. 3 is a schematic circuit diagram of the SiPM according to the embodiment. FIG. 4 is a diagram illustrating an example of a list of count information according to an embodiment. FIG. 5 is a diagram for explaining an example of a time series list of coincidence counting information in the embodiment. FIG. 6 is a flowchart showing an example of a flow of processing executed by the PET device according to the embodiment. FIG. 7 is a diagram for explaining an example of the processes in steps S101 and S102 shown in FIG. FIG. 8 is a diagram for explaining an example of the processes in steps S101 and S102 shown in FIG. FIG. 9 is a diagram for explaining an example of the processes in steps S101 and S102 shown in FIG. FIG. 10 is a graph showing an example of the relationship between the reverse bias voltage and the time resolution of the SiPM according to the embodiment. FIG. 11 is a diagram for explaining an example of a process executed by the PET device according to the embodiment. FIG. 12 is a diagram for explaining an example of a process executed by the PET device according to the embodiment. FIG. 13 is a diagram for explaining an example of a process executed by the PET device according to the embodiment.

Below, an embodiment of a nuclear medicine diagnosis device will be described in detail with reference to the attached drawings. Furthermore, the nuclear medicine diagnosis device according to the present application is not limited to the embodiment shown below. Furthermore, the embodiment can be combined with other embodiments, other modified examples, and conventional techniques to the extent that no contradiction occurs in the content. Furthermore, in the following description, similar components are given common reference numerals, and duplicated descriptions may be omitted.

(Embodiment)
First, a configuration of a nuclear medicine diagnosis apparatus according to an embodiment will be described. In the embodiment, a PET apparatus will be described as an example of the nuclear medicine diagnosis apparatus.

FIG. 1 is a block diagram showing an example of the configuration of a PET device 100 according to an embodiment. As shown in FIG. 1, the PET device 100 according to an embodiment includes a gantry device 10 and a console device 20. The gantry device 10 and the console device 20 are connected by wire or wirelessly and can communicate with each other.

The gantry device 10 detects a pair of gamma rays (annihilation gamma rays) emitted when a positron emitted in the subject P combines with an electron and annihilates, using a semiconductor gamma ray detector (semiconductor detector) 14, and collects counting information by generating counting information from the output signal output from the semiconductor gamma ray detector 14. The subject P has been administered, for example, a radioactive drug (radiopharmaceutical drug) labeled with a positron-emitting nuclide. Gamma rays are an example of radiation.

As shown in FIG. 1, the gantry 10 includes a tabletop 11, a bed 12, a bed driver 13, a semiconductor gamma ray detector 14, a counting information collection circuit 15, a processing circuit 16, a power supply control circuit 17, and a cooling device (cooling module) 18. As shown in FIG. 1, the gantry 10 has a cavity that serves as the imaging port.

The top board 11 is a bed on which the subject P is placed, and is placed on the bed 12. The bed driver 13 moves the top board 11 under the control of the bed control circuit 23 described later. For example, the bed driver 13 moves the top board 11 to move the subject P into the imaging aperture of the gantry device 10.

The semiconductor gamma ray detector 14 counts and detects gamma rays (annihilation gamma rays) emitted from within the subject P. In other words, the semiconductor gamma ray detector 14 counts incident gamma rays (incident gamma rays). As shown in FIG. 1, the semiconductor gamma ray detector 14 has multiple detector modules 14a arranged in a ring shape surrounding the subject P. The detector modules 14a convert the gamma rays emitted from within the subject P into light, and then convert the converted light into an electrical signal.

Here, an example of the configuration of the detector module 14a will be described. FIG. 2 is a diagram showing an example of the configuration of the detector module 14a according to the embodiment. As shown in FIG. 2, the detector module 14a has a scintillator array 14b and a SiPM array 14c.

The scintillator array 14b has multiple scintillators 14d arranged two-dimensionally. The scintillators 14d convert gamma rays emitted from the internal tissue of the subject P into light (scintillation light) that has a peak in the ultraviolet region and output it.

The SiPM array 14c has a plurality of SiPMs 14e arranged two-dimensionally. Each of the plurality of SiPMs 14e is optically connected to each of the plurality of scintillators 14d. The SiPMs 14e convert the scintillation light output from the optically connected scintillators 14d into an electrical signal and output the electrical signal. The SiPMs 14e multiply the scintillation light as an electrical signal by a predetermined gain factor. A pair of one scintillator 14d and one SiPM 14e optically connected to this scintillator 14d corresponds to one detector pixel.

Figure 3 is a schematic circuit diagram of a SiPM 14e according to an embodiment. As shown in the example of Figure 3, each SiPM 14e has a plurality of APD (Avalanche Photo Diode) cells 14f arranged two-dimensionally. The APD cells 14f are photoelectric conversion elements and are silicon-based semiconductors. The APD cells 14f explosively generate electron-hole pairs upon receiving scintillation light from the scintillator 14d. This phenomenon is called an electron avalanche. The electron avalanche causes the APD cells 14f to generate a current pulse.

The multiple APD cells 14f included in each SiPM 14e are connected to the waveform shaping circuit 14h via a common output circuit 14g. The output circuit 14g superimposes the current pulses generated by the multiple APD cells 14f to generate an energy signal, which is a current pulse corresponding to the energy of the gamma ray incident on one detector pixel described above. The output circuit 14g then outputs the generated energy signal to the waveform shaping circuit 14h. The waveform shaping circuit 14h shapes the waveform of the energy signal from the output circuit 14g, and outputs the shaped energy signal to the counting information collection circuit 15. The energy signal output from the waveform shaping circuit 14h in this way is the output signal of the semiconductor gamma ray detector 14.

The counting information collection circuit 15 is an electrical circuit that generates counting information from the output signal of the semiconductor gamma ray detector 14 and stores the generated counting information in the memory 24 described below.

For example, the counting information collecting circuit 15 collects counting information by generating counting information from the output signal of the semiconductor gamma ray detector 14. This counting information includes the detection position, energy value, and detection time of the gamma ray. For example, the counting information includes the scintillator number (P), energy value (E), and detection time (T). The scintillator number (P) is a number that identifies the scintillator 14d that converted the gamma ray into scintillation light. Note that, instead of the scintillator number (P), information indicating the spatial position where the gamma ray was converted into scintillation light in the scintillator 14d may be used. The energy value (E) is the energy value of the gamma ray. The detection time (T) is the time when the gamma ray was converted into scintillation light.

The counting information collection circuit 15 identifies the scintillator number (P) and detection time (T) from the output signal of the semiconductor gamma ray detector 14, and calculates the energy value (E). The counting information collection circuit 15 then generates counting information including the identified scintillator number (P) and detection time (T) and the calculated energy value (E).

Here, the counting information collection circuit 15 calculates the energy value (E) using a time-over-threshold (TOT) method. For example, the counting information collection circuit 15 measures the time at which the output signal from the semiconductor gamma ray detector 14 exceeds a threshold using the time-over-threshold method, and calculates the energy value (E) based on the measured time. The threshold used in such a time-over-threshold method is an example of a parameter value used when calculating the energy of gamma rays.

The counting information collection circuit 15 also determines the actual time at which the gamma rays are converted into scintillation light, and corrects the determined actual time using a known time offset and a known time walk. The counting information collection circuit 15 then determines the corrected time as the detection time (T). Such a time offset and time walk are examples of parameter values used when correcting the detection timing of gamma rays.

Although not shown in FIG. 1, the multiple detector modules 14a are divided into multiple blocks, and each block is provided with a counting information collection circuit 15. For example, when one detector module 14a belongs to one block, the counting information collection circuit 15 is provided for each detector module 14a.

The processing circuit 16 includes an acquisition function 16a and a control function 16b. The acquisition function 16a acquires information related to the count rate of gamma rays incident on the semiconductor gamma ray detector 14. The control function 16b controls the value of the reverse bias voltage (reverse bias voltage value) applied (loaded) to the semiconductor gamma ray detector 14 based on the acquired information related to the count rate of gamma rays. Details of the acquisition function 16a and the control function 16b will be described later.

Here, for example, each processing function of the acquisition function 16a and the control function 16b, which are components of the processing circuit 16 shown in FIG. 1, is recorded in the memory provided in the processing circuit 16 in the form of a program executable by a computer. The processing circuit 16 reads each program from the memory and executes each read program to realize the function corresponding to each program. In other words, the processing circuit 16 in a state in which each program has been read has each function shown in the processing circuit 16 in FIG. 1. The processing circuit 16 is realized by, for example, a processor.

All of the processing functions of the acquisition function 16a and the control function 16b may be recorded in memory in the form of a single program executable by a computer. In this case, the processing circuit 16 reads the program from the memory and executes the read program to realize the acquisition function 16a and the control function 16b corresponding to the program. The acquisition function 16a is an example of an acquisition unit. The control function 16b is an example of a control unit.

The power supply control circuit 17 applies (loads) a reverse bias voltage to the SiPM 14e. For example, under control of the control function 16b of the processing circuit 16, the power supply control circuit 17 applies the reverse bias voltage to the SiPM 14e by supplying power specified by the control function 16b to the SiPM 14e based on power supplied from a power supply (not shown). Here, the value of the reverse bias voltage applied to the SiPM 14e (reverse bias voltage value) is controlled by the control function 16b. The reverse bias voltage is an example of a bias voltage, and the reverse bias voltage value is an example of a bias voltage value.

The processing circuit 16 and the power supply control circuit 17 may be located within the semiconductor gamma ray detector 14. In other words, the semiconductor gamma ray detector 14 may be equipped with the processing circuit 16 and the power supply control circuit 17.

The cooling device 18 cools a part of the PET device 100. For example, the cooling device 18 cools the semiconductor gamma ray detector 14 under the control of the control function 16b. Here, the degree of cooling (cooling level) by the cooling device 18 is controlled by the control function 16b.

The console device 20 accepts operations of the PET device 100 by the operator, controls the capture of PET image data, and reconstructs the PET image data using the counting information collected by the gantry device 10. As shown in FIG. 1, the console device 20 includes an input interface 21, a display 22, a bed control circuit 23, a memory 24, a coincidence counting information generating circuit 25, an image reconstruction circuit 26, and a system control circuit 27. The circuits included in the console device 20 are connected via a bus.

The input interface 21 is realized by a trackball, a switch, a dial, a foot switch, a joystick, etc. The input interface 21 is used by the operator of the PET device 100 to input various instructions and settings, and transfers the input instructions and settings to the system control circuit 27. For example, the input interface 21 is used to input an instruction to start imaging.

Under the control of the system control circuit 27, the display 22 displays a PET image based on the PET image data, and displays a GUI (Graphical User Interface) for receiving various instructions and settings from the operator. The display 22 is an example of a display unit. The display 22 is realized by a liquid crystal monitor, a CRT (Cathode Ray Tube) monitor, or the like. The bed control circuit 23 is an electric circuit that controls the bed driver 13.

The memory 24 stores various data used in the PET device 100. The memory 24 is realized, for example, by a semiconductor memory element such as a random access memory (RAM) or a flash memory, a hard disk, an optical disk, etc.

The memory 24 stores a list of counting information generated by each counting information collection circuit 15. The list of counting information stored in the memory 24 is used for processing by the coincidence counting information generation circuit 25. The list of counting information stored in the memory 24 may be deleted after being used for processing by the coincidence counting information generation circuit 25, or may be stored for a predetermined period of time.

Figure 4 is a diagram for explaining an example of a list of counting information in an embodiment. As shown in Figure 4, the memory 24 stores counting information including a scintillator number (P), an energy value (E), and a detection time (T) in association with a module ID that identifies the detector module 14a.

The memory 24 also stores a time series list of coincidence counting information generated by the coincidence counting information generating circuit 25. The time series list of coincidence counting information stored in the memory 24 is used for processing by the image reconstruction circuit 26. The time series list of coincidence counting information stored in the memory 24 may be deleted after being used for processing by the image reconstruction circuit 26, or may be stored for a predetermined period of time.

Figure 5 is a diagram for explaining an example of a time series list of coincidence counting information in an embodiment. As shown in Figure 5, memory 24 stores sets of counting information in association with coincidence numbers, which are serial numbers of the coincidence counting information. Note that the time series list of coincidence counting information is arranged in approximately chronological order based on the detection times (T) of the counting information.

The memory 24 also stores the PET image data reconstructed by the image reconstruction circuit 26. The PET image based on the PET image data stored in the memory 24 is displayed on the display 22 by the system control circuit 27.

Returning to FIG. 1, the coincidence information generating circuit 25 generates a chronological list of coincidence information using the list of count information generated by the count information collecting circuit 15. For example, the coincidence information generating circuit 25 searches the list of count information stored in the memory 24 for a set of count information in which a pair of gamma rays are counted approximately simultaneously, based on the detection time (T) of the count information. In addition, the coincidence information generating circuit 25 generates coincidence information for each set of searched count information, and stores the generated coincidence information in the memory 24 while arranging it in approximately chronological order.

For example, the coincidence counting information generating circuit 25 generates the coincidence counting information based on the conditions (coincidence counting information generating conditions) for generating the coincidence counting information input by the operator. A time window (time window width) is specified as the coincidence counting information generating condition. For example, the coincidence counting information generating circuit 25 generates the coincidence counting information based on the time window.

For example, the coincidence information generating circuit 25 refers to the memory 24 and searches for a set of counting information whose detection time (T) difference is within the time window between the detector modules 14a. For example, when the coincidence information generating circuit 25 searches for a set of "P11, E11, T11" and "P22, E22, T22" as a set that satisfies the coincidence information generating condition, the coincidence information generating circuit 25 generates the set as coincidence information and stores it in the memory 24. The coincidence information generating circuit 25 may generate the coincidence information using an energy window (energy window width) together with the time window. For example, the coincidence information generating circuit 25 refers to the memory 24 and searches for a set of counting information whose detection time (T) difference is within the time window and whose energy value (E) is within the energy window between the detector modules 14a. The width of the energy window is an example of a parameter value used when calculating the energy of gamma rays. The coincidence information generating circuit 25 may also be provided in the gantry device 10. The coincidence counting information generating circuit 25 is realized, for example, by a processor.

The image reconstruction circuit 26 reconstructs the PET image data. For example, the image reconstruction circuit 26 reads out a time series list of coincidence counting information stored in the memory 24, and reconstructs the PET image data using the read out time series list. The image reconstruction circuit 26 also stores the reconstructed PET image data in the memory 24. The image reconstruction circuit 26 is realized, for example, by a processor.

The system control circuit 27 performs overall control of the PET device 100 by controlling the gantry device 10 and the console device 20. For example, the system control circuit 27 controls imaging in the PET device 100. The system control circuit 27 is realized, for example, by a processor.

The overall configuration of the PET device 100 according to the embodiment has been described above. Here, we will explain the case where, in imaging using the PET device, the value of the reverse bias voltage applied to the SiPM is constant regardless of the gamma ray count rate. In this case, it is not possible to make the most of the performance of the SiPM at low or high count rates. Furthermore, when attempting to make the most of the performance of the SiPM, there is a design constraint that must be met so that the performance of the detector equipped with the SiPM is less likely to deteriorate even at high count rates.

Therefore, based on the above-mentioned configuration, the PET device 100 is configured to achieve both the inherent performance and stability of the semiconductor gamma ray detector 14, as described below.

FIG. 6 is a flowchart showing an example of a flow of a process executed by the PET device 100 according to the embodiment. The process shown in FIG. 6 is executed, for example, when the PET device 100 starts capturing PET image data of the subject P. At the start of the capture, the value of the reverse bias voltage applied to the SiPM 14e is the default value "V2", and the level of cooling by the cooling device 18 is the default cooling level "FR2". At the start of the capture, the time offset is the default value "T _OFFSET 2", and the time walk is the default value "T _WALK 2". At the start of the capture, the width of the energy window is the default value "W2", and the threshold used in the time over threshold method is the default value "Th2".

As shown in FIG. 6, the acquisition function 16a acquires information regarding the count rate of gamma rays incident on the semiconductor gamma ray detector 14 (step S101). Then, the control function 16b determines whether the count rate of the incident gamma rays is a high count rate based on the information regarding the count rate of the incident gamma rays (step S102).

Here, specific examples of information related to the count rate of the incident gamma rays will be described below. Figures 7 to 9 are diagrams for explaining an example of the processing of steps S101 and S102 shown in Figure 6. For example, in step S101, the acquisition function 16a acquires the value of the current flowing through the semiconductor gamma ray detector 14 as information related to the count rate of the incident gamma rays, as shown in Figure 7. More specifically, the acquisition function 16a acquires the value of the current flowing through the SiPM 14e as the value of the current flowing through the semiconductor gamma ray detector 14.

For example, a current flows between the power supply control circuit 17 and the SiPM 14e. It is considered that the larger the current value, the higher the count rate. Therefore, a current measuring device (not shown) measures the value of the current input to the SiPM 14e or the value of the current output from the SiPM 14e. The acquisition function 16a then acquires the value of the current measured by the current measuring device as the value of the current flowing through the semiconductor gamma ray detector 14.

Then, in step S102, the control function 16b determines whether or not the current value acquired by the acquisition function 16a exceeds a threshold value, as shown in FIG. 7. If the current value exceeds the threshold value (step S102: Yes), the control function 16b determines that the count rate of the incident gamma rays is a high count rate. If the current value does not exceed the threshold value (step S102: No), the control function 16b determines that the count rate of the incident gamma rays is not a high count rate.

In addition, when capturing PET image data of a subject P who has been administered a specific type of radiopharmaceutical, which results in a high count rate equal to or greater than a predetermined value, it is considered that the count rate will be a high count rate equal to or greater than a predetermined value. Therefore, as shown in FIG. 8, the acquisition function 16a may acquire information on the type of radiopharmaceutical administered to the subject P to be imaged in step S101. For example, the memory 24 stores examination information. The examination information includes the type of radiopharmaceutical administered to the subject P. Therefore, the acquisition function 16a may acquire the type of radiopharmaceutical administered to the subject P to be imaged from the examination information stored in the memory 24 in step S101.

Then, in step S102, the control function 16b determines whether the type of radiopharmaceutical acquired by the acquisition function 16a is a specific type that results in a high count rate, as shown in FIG. 8. If the type of radiopharmaceutical is a specific type (step S102: Yes), the control function 16b determines that the count rate of the incident gamma rays is a high count rate. If the type of radiopharmaceutical is not a specific type (step S102: No), the control function 16b determines that the count rate of the incident gamma rays is not a high count rate.

In addition, if the part of the subject P to be imaged, to which a radiopharmaceutical has been administered, is a part such as the heart where the radiopharmaceutical is likely to accumulate, the count rate is considered to be a high count rate equal to or higher than a predetermined value. Therefore, as shown in FIG. 9, the acquisition function 16a may acquire information on the part of the subject P to be imaged, to which a radiopharmaceutical has been administered, in step S101. For example, the examination information stored in the memory 24 includes the part of the subject to be imaged. Therefore, the acquisition function 16a may acquire the part of the subject to be imaged from the examination information stored in the memory 24 in step S101. The acquisition function 16a may extract the part of the subject to be imaged from the scout image data or the PET image data.

Then, in step S102, the control function 16b determines whether the part of the subject to be imaged acquired or extracted by the acquisition function 16a is a specific part that has a high count rate, as shown in FIG. 9. If the part of the subject to be imaged is a specific part (step S102: Yes), the control function 16b determines that the count rate of the incident gamma rays is a high count rate. If the part of the subject to be imaged is not a specific part (step S102: No), the control function 16b determines that the count rate of the incident gamma rays is not a high count rate.

The acquisition function 16a may acquire dose information indicating the dose of gamma rays in step S101. In this case, the control function 16b may determine whether the count rate of the incident gamma rays is a high count rate by determining whether the dose indicated by the dose information exceeds a threshold in step S102. Specifically, when the dose indicated by the dose information exceeds the threshold (step S102: Yes), the control function 16b determines that the count rate of the incident gamma rays is a high count rate. When the dose indicated by the dose information does not exceed the threshold (step S102: No), the control function 16b determines that the count rate of the incident gamma rays is not a high count rate.

The acquisition function 16a may also acquire the value of the reverse bias voltage in step S101. In this case, the control function 16b may determine whether the count rate of the incident gamma rays is a high count rate by determining whether the value of the reverse bias voltage exceeds a threshold in step S102. Specifically, when the value of the reverse bias voltage exceeds the threshold (step S102: Yes), the control function 16b determines that the count rate of the incident gamma rays is a high count rate. When the value of the reverse bias voltage does not exceed the threshold (step S102: No), the control function 16b determines that the count rate of the incident gamma rays is not a high count rate.

In addition, in step S101, the acquisition function 16a may acquire information indicating whether or not the count rate is high, which is input by the user via the input interface 21. In this case, in step S102, the control function 16b may determine whether or not the count rate of the incident gamma rays is a high count rate by determining whether or not the information acquired in step S101 indicates a high count rate.

Specifically, if the information acquired in step S101 indicates a high count rate (step S102: Yes), control function 16b determines that the count rate of the incident gamma rays is a high count rate. In addition, if the information acquired in step S101 does not indicate a high count rate (step S102: No), control function 16b determines that the count rate of the incident gamma rays is not a high count rate.

Returning to the explanation of FIG. 6, if the count rate of the incident gamma rays is not a high count rate (step S102: No), the control function 16b maintains the various default values and default cooling levels described above without changing them (step S103). Then, the control function 16b ends the process shown in FIG. 6.

Figure 10 is a graph showing an example of the relationship between the reverse bias voltage and the time resolution of the SiPM 14e according to the embodiment. In the graph shown in Figure 10, the horizontal axis indicates the value of the reverse bias voltage, and the vertical axis indicates an index indicating the time resolution of the SiPM 14e. In the example of Figure 10, the smaller the index indicating the time resolution of the SiPM 14e, the higher the time resolution. As shown in the example of Figure 10, when the gamma ray count rate is not a high count rate, "V2" is used as the default value of the reverse bias voltage applied to the SiPM 14e so that the time resolution of the SiPM 14e is maximized (highest) (i.e., so that the index is at the minimum value "TR2").

On the other hand, if the count rate of the incident gamma rays is a high count rate (step S102: Yes), the control function 16b controls the power supply control circuit 17 to apply a reverse bias voltage of "V1" smaller than the default value "V2" to the SiPM 14e in order to reduce the value of the current flowing through the SiPM 14e (step S104). As a result, the power supply control circuit 17 applies a reverse bias voltage of "V1" to the SiPM 14e. As shown in the example of FIG. 10, when the value of the reverse bias voltage is "V1", the index indicating the time resolution of the SiPM 14e becomes "TR1" larger than "TR2". Thus, in step S104, the control function 16b controls the reverse bias voltage value applied to the semiconductor gamma ray detector 14 based on information regarding the count rate of the incident gamma rays.

Here, when the gamma ray count rate is high, the power consumption of the semiconductor gamma ray detector 14 increases and the amount of heat generated increases. Therefore, the control function 16b controls the cooling device 18 so that the level of cooling of the semiconductor gamma ray detector 14 by the cooling device 18 becomes "FR1", which is higher than the default cooling level "FR2" (step S105).

Then, the control function 16b controls various parameters, etc. based on the value "V1" of the reverse bias voltage changed in step S104 (step S106). Then, the control function 16b ends the process shown in FIG. 6.

The processing of step S106 will be described with a specific example with reference to Figures 11 to 13. Figures 11 to 13 are diagrams for explaining an example of the processing executed by the PET device 100 according to the embodiment.

For example, the dark count rate is changed according to the value of the reverse bias voltage. The dark count rate is, for example, a count rate based on the dark current flowing through the SiPM14e. For example, as the value of the reverse bias voltage increases, the dark count rate also increases, and as the value of the reverse bias voltage decreases, the dark count rate also decreases. The threshold value used in the above-mentioned time-over-threshold method is used to distinguish between the count rate based on the incident gamma rays and the dark count rate. In other words, such a threshold value is used to count the gamma rays that are actually incident, without counting the gamma rays based on the dark current. Here, since the value of the reverse bias voltage is changed to "V1" in step S104, it is preferable to change the threshold value used in the time-over-threshold method.

Therefore, in step S106, the control function 16b changes the threshold value used in the threshold exceeding time method as shown in FIG. 11. For example, the memory of the processing circuit 16 stores a correspondence between each of a plurality of reverse bias voltages and an optimal threshold value corresponding to each reverse bias voltage. Here, an example of the optimal threshold value will be described. For example, the optimal threshold value refers to a threshold value at which the probability that the count is not a dark count (a count based on the dark current flowing through the SiPM 14e) but is a count due to an actually incident gamma ray (gamma ray event) is statistically higher than a predetermined probability. For example, when the dark count rate increases due to an increase in thermal noise or an increase in the reverse bias voltage, the optimal threshold value becomes higher. Then, in step S106, the control function 16b refers to the contents stored in the memory and specifies the optimal threshold value "Th1" corresponding to the value "V1" of the reverse bias voltage. Then, in step S106, the control function 16b controls the counting information collection circuit 15 to change the threshold value used in the threshold exceeding time method from the default value "Th2" to the threshold value "Th1". For example, the threshold value "Th1" is smaller than the default value "Th2." In response to this control, the counting information collection circuit 15 changes the threshold value used in the time-over-threshold method from the default value "Th2" to the threshold value "Th1." In this way, the control function 16b controls the threshold value based on the reverse bias voltage value controlled in step S104.

The time offset and time walk described above are also affected by the change in the value of the reverse bias voltage. Specifically, when the value of the reverse bias voltage is changed, the carrier movement speed of the SiPM 14e is also changed, so the time offset and time walk are also affected. Therefore, in step S106, the control function 16b changes the time offset and time walk as shown in FIG. 12. For example, the memory of the processing circuit 16 stores each of a plurality of reverse bias voltages and an optimal time offset and an optimal time walk corresponding to each reverse bias voltage in association with each other. Here, an example of the optimal time offset and the optimal time walk will be described. For example, the optimal time offset refers to a time offset set so that the reference time (for example, the time when the gamma rays are estimated to have interacted, excluding the signal processing delay) is the same for all channels. Also, for example, the optimal time walk refers to a time walk that appropriately corrects the relationship between the energy of the gamma rays and the rise time when the gain is changed. Then, in step S106, the control function 16b refers to the stored contents of the memory, and specifies the optimum time offset "T _OFFSET 1" and the optimum time walk "T _WALK 1" that correspond to the reverse bias voltage value "V1".

Then, in step S106, the control function 16b controls the counting information collecting circuit 15 to change the time offset from the default value "T _OFFSET 2" to "T _OFFSET 1". Also, in step S106, the control function 16b controls the counting information collecting circuit 15 to change the time walk from the default value "T _WALK 2" to "T _WALK 1". In response to these controls, the counting information collecting circuit 15 changes the time offset from the default value "T _OFFSET 2" to "T _OFFSET 1" and changes the time walk from the default value "T _WALK 2" to "T _WALK 1". In this way, the control function 16b controls the time offset and the time walk based on the reverse bias voltage value controlled in step S104.

The energy window described above is also affected by the change in the value of the reverse bias voltage. Therefore, in step S106, the control function 16b changes the width and position of the energy window as shown in FIG. 13. For example, the memory of the processing circuit 16 stores a correspondence between each of a plurality of reverse bias voltages and the width of the optimal energy window corresponding to each reverse bias voltage. Here, an example of the width of the optimal energy window will be described. For example, the width of the optimal energy window refers to an energy window that is adjusted so that the total count number does not change even if the energy spectrum shifts or shrinks. Then, in step S106, the control function 16b refers to the contents stored in the memory and identifies the width "W1" of the optimal energy window corresponding to the value "V1" of the reverse bias voltage.

Then, in step S106, the control function 16b notifies the coincidence information generation circuit 25 to change the width of the energy window from the default value "W2" to "W1" as shown in FIG. 13. Specifically, the control function 16b adds the changed energy window width "W1" to the list of counting information, and stores the list of counting information to which the changed energy window width "W1" has been added in the memory 24. The coincidence information generation circuit 25 acquires the energy window width "W1" added to the list of counting information. Then, the coincidence information generation circuit 25 changes the width of the energy window from the default value "W2" to the acquired "W1". Thus, in step S106, the control function 16b controls the width of the energy window based on the reverse bias voltage value controlled in step S104.

Note that multiple files may be stored in the memory 24, and multiple energy window widths may be registered in each of the multiple files. The coincidence information generation circuit 25 may set the energy window width registered in one of the multiple files as the energy window width used when generating a time series list of coincidence information. In this case, in step S106, the control function 16b adds the identifier of the file in which the changed energy window width "W1" is registered to the list of counting information. The control function 16b then stores the list of counting information to which the file identifier is added in the memory 24. The coincidence information generation circuit 25 then acquires the identifier of the file added to the list of counting information. The coincidence information generation circuit 25 then acquires the energy window width "W1" registered in the file indicated by the acquired identifier among the multiple files stored in the memory 24. The coincidence information generation circuit 25 then changes the energy window width from the default value "W2" to the acquired "W1".

The above describes the PET device 100 according to the embodiment. According to the embodiment, in the case of a high count rate, the reverse bias voltage in the semiconductor gamma ray detector 14 is changed according to the count rate, and various parameters and cooling levels are changed according to the changed reverse bias voltage, thereby making it possible to achieve both the inherent performance and stability of the semiconductor gamma ray detector 14 equipped with the SiPM 14e.

Furthermore, according to the embodiment, not only can the performance of the semiconductor gamma ray detector 14 equipped with the SiPM 14e be safely maximized even under different count rate conditions, but the downstream circuitry including the power supply system can be used as is for SiPMs with different specifications.

In addition, according to the above-described embodiment, in a clinical examination, if the subject P is imaged again after a period of time has elapsed since the concentration of the radioactive drug in the subject P is low, a negative determination is made in step S102 of FIG. 6. Therefore, when the subject P is imaged again, the imaging is performed with the default settings.

In the first embodiment, the case has been described where the PET device 100 changes the value of the reverse bias voltage in two stages depending on the count rate. However, the PET device 100 may change the value of the reverse bias voltage in more than two stages depending on the count rate. In addition, the PET device 100 may continuously change the value of the reverse bias voltage using a predetermined formula in accordance with the characteristics of the SiPM 14e depending on the count rate.

The term "processor" used in the above description means a circuit such as a CPU (Central Processing Unit), a GPU (Graphics Processing Unit), an Application Specific Integrated Circuit (ASIC), or a programmable logic device (e.g., a Simple Programmable Logic Device (SPLD), a Complex Programmable Logic Device (CPLD), or a Field Programmable Gate Array (FPGA)). The processor realizes its function by reading and executing a program stored in a memory in the processor circuit. Instead of storing a program in a memory in the processor circuit, the program may be stored in a memory provided separately from the processor circuit. In this case, the processor realizes its function by reading and executing the program stored in this memory. Note that each processor in this embodiment is not limited to being configured as a single circuit for each processor, but may be configured as a single processor by combining multiple independent circuits to realize its function.

According to at least one of the embodiments of the PET device 100 described above, it is possible to achieve both the inherent performance and stability of the semiconductor gamma ray detector 14.

Although several embodiments have been described, these embodiments are presented as examples and are not intended to limit the scope of the invention. These embodiments can be implemented in various other forms, and various omissions, substitutions, modifications, and combinations of embodiments can be made without departing from the spirit of the invention. These embodiments and their modifications are within the scope of the invention and its equivalents as set forth in the claims, as well as the scope and spirit of the invention.

14 Semiconductor gamma ray detector 16a Acquisition function 16b Control function

Claims (5)

a semiconductor gamma ray detector for counting incident gamma rays, the semiconductor gamma ray detector including a scintillator for converting the incident gamma rays into light, and a photomultiplier for converting the light into an electrical signal and outputting the electrical signal;
an acquisition unit that acquires a value of a current flowing through the semiconductor gamma ray detector as information regarding a gamma ray count rate;
a control unit that controls a bias voltage value applied to the photomultiplier device of the semiconductor gamma ray detector based on information about the count rate;
A nuclear medicine diagnostic device comprising: a semiconductor gamma ray detector for counting incident gamma rays, the semiconductor gamma ray detector including a scintillator for converting the incident gamma rays into light, and a photomultiplier for converting the light into an electrical signal and outputting the electrical signal;
an acquisition unit that acquires information regarding a part of a subject to be imaged as information regarding a gamma ray count rate;
a control unit that controls a bias voltage value applied to the photomultiplier device of the semiconductor gamma ray detector based on information about the count rate;
A nuclear medicine diagnostic device comprising: a semiconductor gamma ray detector for counting incident gamma rays, the semiconductor gamma ray detector including a scintillator for converting the incident gamma rays into light, and a photomultiplier for converting the light into an electrical signal and outputting the electrical signal;
an acquisition unit that acquires information regarding a type of drug administered to a subject to be imaged as information regarding the gamma ray count rate;
a control unit that controls a bias voltage value applied to the photomultiplier device of the semiconductor gamma ray detector based on information about the count rate;
A nuclear medicine diagnostic device comprising: a semiconductor gamma ray detector for counting incident gamma rays, the semiconductor gamma ray detector including a scintillator for converting the incident gamma rays into light, and a photomultiplier for converting the light into an electrical signal and outputting the electrical signal;
an acquisition unit for acquiring information regarding a gamma ray count rate;
a control unit that controls a bias voltage value applied to the photomultiplier device of the semiconductor gamma ray detector based on information about the count rate;
Equipped with
The control unit controls a value of a parameter used in calculating the energy of the gamma ray based on the controlled bias voltage value.
Nuclear medicine diagnostic equipment. a semiconductor gamma ray detector for counting incident gamma rays, the semiconductor gamma ray detector including a scintillator for converting the incident gamma rays into light, and a photomultiplier for converting the light into an electrical signal and outputting the electrical signal;
an acquisition unit for acquiring information regarding a gamma ray count rate;
a control unit that controls a bias voltage value applied to the photomultiplier device of the semiconductor gamma ray detector based on information about the count rate;
Equipped with
the control unit controls a value of a parameter used when correcting the detection timing of the gamma ray based on the controlled bias voltage value.
Nuclear medicine diagnostic equipment.

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