JP6878131B2 - Medical diagnostic imaging equipment - Google Patents

Description

An embodiment of the present invention relates to a medical diagnostic imaging apparatus.

Conventionally, as a medical image diagnostic device capable of performing functional diagnosis in a living tissue of a subject, a positron emission tomography (PET) device and a single photon emission tomography device (SPECT (Single Photon)) have been used. Emission Computed Tomography) equipment) and the like are known.

The detector module, which is a gamma ray detector in a PET device, receives scintillation photons emitted when gamma rays emitted from a subject enter a scintillator with a photomultiplier tube (PMT) and converts them into electrical signals. .. For example, one scintillation photon that collides with the photoelectric surface of a photomultiplier tube ejects one electron by the photoelectric effect. A plurality of positively charged dynodes are provided after the photoelectric surface. Electrons accelerated by an electrically attractive force travel toward the dynodes provided in a plurality of stages, and the electrons collide with the dynodes. When an electron collides with a dynode, millions of electrons are output from the photomultiplier tube as an electrical signal. That is, one electron ejected from one scintillation photon is amplified into several million electrons and flows as an electric signal. The amplification factor of the number of electrons at this time is called the gain factor. This gain rate varies from photomultiplier tube to photomultiplier tube. For example, the gain rate varies several times.

Therefore, even if the same number of scintillation photons are incident on the photomultiplier tube, the output current will be different for each photomultiplier tube. Therefore, since the number of scintillation photons represents the energy of the incident gamma ray, in the detector module composed of a plurality of photomultiplier tubes, calibration for aligning the outputs of the photomultiplier tubes by the amplifier circuit in the subsequent stage, that is, energy calibration is performed. There is a need to do.

^{Therefore, using a sealed dot source such as 68} Ge or a sealed linear source that emits monochromatic gamma rays with less scattering, gamma rays from the sealed dot source or sealed linear source are used as a detector module. It is common to generate a histogram of the time integration value of the output signal for each photomultiplier tube, and adjust the amplification factor in the amplifier circuit so that the peak positions are approximately the same for all photomultiplier tubes. It is done in.

Japanese Unexamined Patent Publication No. 62-043584 International Publication No. 2007/043137 Japanese Unexamined Patent Publication No. 2014-176620

An object to be solved by the present invention is to provide a medical image diagnostic apparatus capable of performing energy calibration with high accuracy while suppressing a decrease in operating rate.

The medical image diagnostic apparatus of the embodiment includes a plurality of converters, a first specific unit, a second specific unit, and a correction unit. Multiple transducers output electrical signals based on the incident radiation. The first specific unit is the signal intensity corresponding to the peak of the number of radiations based on the relationship between the signal intensity of the electric signal output from the converter and the number of incident radiations for each of the plurality of converters. The first signal strength is specified. The second specific part is the energy of the incident radiation without scattering for each of the plurality of converters based on the relationship between the signal intensity and the number of radiations at an intensity higher than the first signal intensity. The second signal strength, which is the signal strength corresponding to the above, is specified. The correction unit corrects the signal strength of the electric signal output from each of the plurality of converters so that the second signal strength specified for each of the plurality of converters matches the target signal strength.

FIG. 1 is a diagram showing an example of the configuration of the PET device according to the first embodiment. FIG. 2 is a diagram for explaining a list of counting information according to the first embodiment. FIG. 3 is a diagram for explaining a time-series list of coincidence counting information according to the first embodiment. FIG. 4 is a diagram showing an example of the configuration of the detector module according to the first embodiment. FIG. 5 is a diagram showing an example of a histogram based on an electric signal obtained by clinical imaging. FIG. 6 is a diagram showing an example of a histogram based on an electric signal obtained by clinical imaging. FIG. 7 is a diagram for explaining an example of processing executed by the first specific function according to the first embodiment. FIG. 8 is a diagram for explaining an example of processing executed by the second specific function according to the first embodiment. FIG. 9 is a diagram for explaining a case where the second specific function according to the first embodiment takes into account the influence of pile-up when specifying the signal strength of the peak. FIG. 10 is a diagram for explaining an example of processing executed by the correction function according to the first embodiment. FIG. 11 is a diagram for explaining an example of processing executed by the correction function according to the first embodiment. FIG. 12 is a flowchart showing an example of a flow of processing executed by the PET apparatus according to the first embodiment. FIG. 13 is a flowchart showing an example of a flow of processing executed by the PET apparatus according to the second embodiment. FIG. 14 is a diagram showing an example of the configuration of the SPECT apparatus according to the third embodiment.

Hereinafter, each embodiment of the medical diagnostic imaging apparatus will be described in detail with reference to the accompanying drawings. In addition, each embodiment can be combined appropriately.

(First Embodiment)
First, the configuration of the nuclear medicine imaging device as the medical image diagnostic device according to the first embodiment will be described. In the first embodiment, a PET apparatus will be described as an example of such a nuclear medicine imaging apparatus.

FIG. 1 is a diagram showing an example of the configuration of the PET device 100 according to the first embodiment. As shown in FIG. 1, the PET device 100 according to the embodiment includes a gantry device 10 and a console device 20.

The gantry device 10 surrounds the subject P in a ring shape with a pair of gamma rays (paired annihilation gamma rays) emitted when the positrons emitted in the subject P combine with electrons and annihilate. The counting information is collected by detecting by the arranged detector module and generating the counting information from the output signal (electrical signal) of the detector module 14. Subject P is administered, for example, a radiopharmaceutical labeled with a positron emitting nuclide. Gamma rays are an example of radiation.

As shown in FIG. 1, the gantry device 10 includes a top plate 11, a sleeper 12, a sleeper drive circuit 13, a plurality of detector modules 14, and a counting information collecting circuit 15. As shown in FIG. 1, the gantry device 10 has a cavity that serves as an image pickup port.

The top plate 11 is a bed on which the subject P is placed, and is arranged on the bed 12. The sleeper drive circuit 13 moves the top plate 11 under the control of the sleeper control circuit 23, which will be described later. For example, the sleeper drive circuit 13 moves the subject P into the image pickup port of the gantry device 10 by moving the top plate 11.

The detector module 14 detects a pair of gamma rays emitted when the positrons emitted in the subject P combine with electrons and annihilate, and outputs an electric signal based on the detected pair of gamma rays. As shown in FIG. 1, a plurality of detector modules 14 are arranged so as to surround the subject P in a ring shape. The detector module 14 converts the gamma rays emitted from the subject P into light, and converts the converted light into an electric signal. The configuration of the detector module 14 will be described later.

The counting information collecting circuit 15 generates counting information from the output signal of the detector module 14, and stores the generated counting information in the data storage circuit 24 described later. For example, the counting information collecting circuit 15 is realized by a processor.

For example, the counting information collecting circuit 15 collects counting information by generating counting information from the output signal of the detector module 14. This counting information includes the gamma ray detection position, the energy value, and the detection time. For example, as will be described later, the counting information includes a scintillator number (P), an energy value (E), and a detection time (T). Although not shown in FIG. 1, the plurality of detector modules 14 are divided into a plurality of blocks, and each block is provided with a counting information collecting circuit 15. For example, when one detector module 14 is one block, a counting information collecting circuit 15 is provided for each detector module 14.

The console device 20 accepts the operation of the PET device 100 by the user, controls the imaging of the PET image, and reconstructs the PET image using the counting information collected by the gantry device 10. As shown in FIG. 1, the console device 20 includes an input circuit 21, a display 22, a sleeper control circuit 23, a data storage circuit 24, a simultaneous counting information generation circuit 25, an image reconstruction circuit 26, and system control. A circuit 27 and a correction circuit 28 are provided. Each circuit or the like included in the console device 20 is connected via a bus.

The input circuit 21 is used by the user of the PET device 100 to input various instructions and various settings. The input circuit 21 transfers various input instructions and various settings to the system control circuit 27. For example, the input circuit 21 is used for inputting an imaging start instruction and an imaging end instruction. For example, the input circuit 21 is realized by a mouse, a keyboard, a button, a panel switch, a touch command screen, a foot switch, a trackball, a joystick, and the like.

The display 22 is realized by a liquid crystal monitor, a CRT (Cathode Ray Tube) monitor, a touch panel, or the like referred to by the user. The display 22 displays a PET image under the control of the system control circuit 27, and displays a GUI (Graphical User Interface) for receiving various instructions and various settings from the user.

The sleeper control circuit 23 controls the sleeper drive circuit 13. For example, the sleeper control circuit 23 is realized by a processor.

The data storage circuit 24 stores various data used in the PET device 100. The data storage circuit 24 is realized by, for example, a semiconductor memory element such as a RAM (Random Access Memory) or a flash memory, a hard disk, an optical disk, or the like.

The data storage circuit 24 stores a list of counting information generated by the counting information collecting circuit 15. The list of counting information stored in the data storage circuit 24 is used for processing by the coincidence counting information generation circuit 25. The list of counting information stored in the data storage circuit 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.

FIG. 2 is a diagram for explaining a list of counting information according to the first embodiment. As shown in FIG. 2, the data storage circuit 24 associates the module ID that identifies the detector module 14 with counting information including a scintillator number (P), an energy value (E), and a detection time (T). Remember.

Further, the data storage circuit 24 stores a time-series list of the coincidence counting information generated by the coincidence counting information generation circuit 25. Further, the time series list of the coincidence counting information stored in the data storage circuit 24 is used for processing by the image reconstruction circuit 26. The time series list of the coincidence counting information stored in the data storage circuit 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.

FIG. 3 is a diagram for explaining a time-series list of coincidence counting information according to the first embodiment. As shown in FIG. 3, the data storage circuit 24 has a coincidence number No. which is a serial number of the coincidence counting information. Stores a set of counting information in association with. The time-series list of the coincidence counting information is arranged in chronological order based on the detection time (T) of the counting information.

Further, the data storage circuit 24 stores the PET image reconstructed by the image reconstruction circuit 26. Further, the PET image stored in the data storage circuit 24 is displayed on the display 22 by the system control circuit 27. In addition, the data storage circuit 24 stores various programs.

Returning to FIG. 1, the coincidence counting information generation circuit 25 generates a time series list of coincidence counting information using the list of counting information generated by the counting information collecting circuit 15. For example, the simultaneous counting information generation circuit 25 searches the list of counting information stored in the data storage circuit 24 for a set of counting information in which a pair of gamma rays are counted substantially simultaneously, based on the detection time (T) of the counting information. To do. Further, the coincidence counting information generation circuit 25 generates the coincidence counting information for each set of the counting information obtained as a result of the search, and stores the generated coincidence counting information in the data storage circuit 24 while arranging them in roughly chronological order. ..

For example, the coincidence counting information generation circuit 25 generates the coincidence counting information based on the conditions for generating the coincidence counting information input by the user (coincidence counting information generation condition). The coincidence counting information generation condition includes the time window width. For example, the coincidence counting information generation circuit 25 generates the coincidence counting information based on the time window width.

For example, the coincidence counting information generation circuit 25 refers to the list of counting information stored in the data storage circuit 24, and sets the counting information in which the time difference of the detection time (T) is within the time window width, the detector module 14 Search between. For example, when the coincidence counting information generation circuit 25 searches for a set of "P11, E11, T11" and "P22, E22, T22" as a set satisfying the coincidence counting information generation condition, this set is generated as the coincidence counting information. Then, it is stored in the data storage circuit 24. The coincidence counting information generation circuit 25 may generate the coincidence counting information by using the energy window width together with the time window width. Further, the coincidence counting information generation circuit 25 may be provided in the gantry device 10.

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

The system control circuit 27 controls the entire 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. For example, the system control circuit 27 is realized by a processor.

The correction circuit 28 includes a first specific function 28a, a second specific function 28b, and a correction function 28c. Here, for example, each processing function of the first specific function 28a, the second specific function 28b, and the correction function 28c, which are the components of the correction circuit 28 shown in FIG. 1, is data in the form of a program that can be executed by a computer. It is recorded in the storage circuit 24. The correction circuit 28 is a processor that realizes a function corresponding to each program by reading each program from the data storage circuit 24 and executing each read program. In other words, the correction circuit 28 in the state where each program is read out has each function shown in the correction circuit 28 of FIG. The first specific function 28a is an example of the first specific unit. The second specific function 28b is an example of the second specific part. The correction function 28c is an example of a correction unit. The first specific function 28a, the second specific function 28b, and the correction function 28c will be described later.

The word "processor" used in the above description is, for example, a CPU (Central Processing Unit), a GPU (Graphics Processing Unit), an integrated circuit for a specific application (Application Specific Integrated Circuit: ASIC), or a programmable logic device (ASIC). For example, it means a circuit such as a simple programmable logic device (SPLD), a complex programmable logic device (CPLD), and a field programmable gate array (FPGA). The processor realizes the function by reading and executing the program stored in the data storage circuit 24. Instead of storing the program in the data storage circuit 24, the program may be directly incorporated in the circuit of the processor. In this case, the processor reads the program embedded in the circuit and executes the read program to realize the function. It should be noted that each processor of the present embodiment is not limited to the case where each processor is configured as a single circuit, and a plurality of independent circuits may be combined to form one processor to realize its function. Good.

The overall configuration of the PET apparatus according to the first embodiment has been described above.

Next, an example of the configuration of the detector module 14 according to the first embodiment will be described. FIG. 4 is a diagram showing an example of the configuration of the detector module 14 according to the first embodiment. As shown in FIG. 4, the detector module 14 is a photon counting type, anger type detector, and includes a plurality of scintillators 141, a plurality of photomultiplier tubes (PMT) 142, and a light guide 143. And a plurality of amplifier circuits 144.

The scintillator 141 converts a pair of gamma rays emitted when the cations emitted in the subject P combine with electrons and annihilate them into scintillation photons (optical photons), and outputs scintillation photons. For example, the scintillator 141 outputs one scintillation photon when one gamma ray is incident. The scintillator 141 is formed by, for example, scintillator crystals such as LaBr3 (Lanthanum Bromide), LYSO (Lutetium Yttrium Oxyorthosilicate), LSO (Lutetium Oxyorthosilicate), and LGSO (Lutetium Gadolinium Oxyorthosilicate). As shown in the example of FIG. 4, the scintillators 141 are arranged two-dimensionally.

The photomultiplier tube 142 multiplies the scintillation photons output from the scintillator 141 and converts them into an electrical signal. As shown in the example of FIG. 4, a plurality of photomultiplier tubes 142 are arranged. The photomultiplier tube 142 has a photocathode that receives scintillation photons and generates photoelectrons, a multi-stage dynode that provides an electric field that accelerates the generated photoelectrons, and an anode that is an electron outflow port. The electrons emitted from the photocathode due to the photoelectric effect are accelerated toward the dynode and collide with the surface of the dynode, knocking out a plurality of electrons. By repeating this phenomenon over multiple stages of dynodes, the number of electrons is multiplied in an avalanche manner, and the number of electrons at the anode reaches several millions. In such an example, the gain rate of the photomultiplier tube 142 is several million times. For example, when one scintillation photon is incident on the photomultiplier tube 142, the photomultiplier tube 142 outputs an electric signal consisting of several million electrons. Further, for amplification utilizing the avalanche phenomenon, a voltage of 1000 volts or more is usually applied between the dynode and the anode. The photomultiplier tube 142 is an example of a converter.

The light guide 143 transmits the scintillation photons output from the scintillator 141 to the photomultiplier tube 142. The light guide 143 is formed of, for example, a plastic material having excellent light transmission.

As described above, the detector module 14 converts the pair annihilation gamma rays emitted from the subject P into scintillation photons by the scintillator 141, and converts the converted scintillation photons into electric signals by the photomultiplier tube 142 and outputs the scintillation photons. .. That is, the detector module 14 is an indirect conversion type detector.

The detector module 14 calculates the position of the scintillator 141 that outputs the light by calculating the center of gravity of the outputs of the plurality of photomultiplier tubes 142 in which the light is incident. The logic for calculating the position of such a scintillator 141 is called, for example, anger logic. When calculating the position of the scintillator 141 by such logic, the number of photomultiplier tubes 142 can be smaller than the number of scintillators 141, but the position of the scintillator 141 is identified from the coordinates obtained as a result of the calculation of the center of gravity. There is a need.

Each of the plurality of amplifier circuits 144 is connected to the subsequent stage of each of the photomultiplier tubes 142. The amplifier circuit 144 amplifies the electric signal output from the photomultiplier tube 142 at a predetermined amplification factor and outputs it to the correction circuit 28.

Here, the gain ratio of the photomultiplier tube 142 described above is unique to each photomultiplier tube 142. Therefore, for example, even when the same number of scintillation photons are incident on the photomultiplier tube 142, the signal intensity of the output electric signal becomes a value peculiar to each photomultiplier tube 142, and all of them. Signal strengths may not be approximately the same. Therefore, when the same number of scintillation photons are incident on each photomultiplier tube 142, the amplification factor of each amplifier circuit 144 is substantially the same so that the signal intensity of the electric signal output from each amplifier circuit 144 is substantially the same. Energy calibration is performed to adjust.

In the above-mentioned energy calibration, the following signal intensities are used as the signal intensities of the electric signals output from each photomultiplier tube 142. For example, in a histogram where the horizontal axis is the signal strength of an electric signal and the vertical axis is the number of incidents of annihilation gamma rays, the signal strength corresponding to the maximum number of events is used. That is, for each photomultiplier tube 142, the signal strength of the peak in the histogram is used as the signal strength of the electrical signal. Then, in the above-mentioned energy calibration, the amplification factor of each amplifier circuit 144 is adjusted so that the peak signal intensities are substantially the same for all the photomultiplier tubes 142.

As described above, in the energy calibration, the signal strength of the peak in the histogram is used for each photomultiplier tube 142. Therefore, to accurately identify the signal strength of this peak is to perform the energy calibration with high accuracy. Leads to.

Here, although details will be described later, when a scattered ray (scattered gamma ray) of a pair annihilation gamma ray is incident on the scintillator 141 in addition to the pair annihilation gamma ray, the peak of the histogram based only on the pair annihilation gamma ray is accurately specified. It becomes difficult to do. In the histogram based only on pair annihilation gamma rays, of the pair annihilation gamma rays and scattered gamma rays, only the pair annihilation gamma rays are incident on the scintillator 141, and the pair annihilation gamma rays are converted into scintillation photons by the scintillator 141. Refers to the histogram obtained from the electric signal when the scintillated photon is converted into an electric signal by. Therefore, it is conceivable to perform the above-mentioned energy calibration using a sealed point-shaped radiation source such as ⁶⁸ Ge or a sealed linear radiation source that emits monochromatic gamma rays with less scattering. In this case, for example, a maintenance period of about one week is provided separately from the period of clinical imaging of the subject P, and energy calibration is performed over a period of about one day during this maintenance period. This energy calibration is performed, for example, once every three months. As described above, in addition to the fact that the energy calibration is not executed for a relatively long period of time, the gain rate of the photomultiplier tube 142 changes with time, so that the performance of the PET apparatus deteriorates for a relatively long period of time. In addition, such a maintenance period is a downtime during which clinical imaging, which is the original purpose of the PET apparatus, cannot be performed. Therefore, providing such a maintenance period is also a factor of lowering the operating rate of the PET apparatus.

Therefore, in the first embodiment, the following processing is performed in order to suppress a decrease in the operating rate of the PET device 100 and a decrease in the performance of the PET device 100 for a long period of time. That is, the PET apparatus 100 according to the first embodiment does not perform energy calibration using a sealed point radiation source during the maintenance period, but uses an electric signal obtained by clinical imaging of the subject P for energy. Calibrate.

However, in addition to the paired annihilation gamma rays, scattered rays of paired annihilation gamma rays are emitted from the subject P during clinical imaging. That is, the subject P is a radiation source accompanied by scattering of photons. Therefore, in clinical imaging, the detector module 14 converts scattered gamma rays into scintillation photons in addition to pair annihilation gamma rays, and converts the scintillation photons into electrical signals for output. Therefore, the histogram based on the pair annihilation gamma rays and the scattered gamma rays emitted from the subject P is a histogram as described below. 5 and 6 are diagrams showing an example of a histogram based on an electric signal obtained by clinical imaging. 5 and 6 show a histogram in which the horizontal axis is the signal intensity of the electric signal and the vertical axis is the number of events, which is the number of paired annihilation gamma rays and scattered gamma rays incident on the scintillator 141. For example, the histogram based on the pair annihilation gamma ray and the scattered gamma ray is obtained as a result of combining the histogram 30 based on the pair annihilation gamma ray shown in FIG. 5 and the histogram 31 based on the scattered gamma ray, and the histogram 32 shown in the example of FIG. Become.

In the histogram based on pair annihilation gamma rays and scattered gamma rays, pair annihilation gamma rays and scattered gamma rays are incident on the scintillator 141, the pair annihilation gamma rays and scattered gamma rays are converted into scintillation photons by the scintillator 141, and scintillation is performed by the photoelectron multiplier tube 142. Refers to a histogram obtained from an electric signal when a photon is converted into an electric signal. Similarly, in the histogram based on scattered gamma rays, of the pair annihilation gamma rays and scattered gamma rays, only the scattered gamma rays are incident on the scintillator 141, the scattered gamma rays are converted into scintillation photons by the scintillator 141, and the scintillation photons are converted by the photoelectron multiplier tube 142. Refers to a histogram obtained from an electric signal when is converted into an electric signal.

In general, gamma rays lose energy when scattered. Therefore, the energy of scattered gamma rays is lower than that of paired annihilation gamma rays. Therefore, as can be seen from the contents shown in FIGS. 5 and 6, the scattered gamma rays affect the shape of the portion of the histogram 30 on the low intensity side from the peak Pt in the histogram 30 based on the pair annihilation gamma rays. For example, as shown in the example of FIG. 6, under the influence of scattered gamma rays, the position of the peak Pt is shifted by x ₁ minute in the horizontal direction to peak Pf. Therefore, in the energy calibration, if the peak Pf of the histogram based on the electrical signal obtained by clinical imaging is simply specified and the signal strength of the peak Pf is used, the accuracy of the energy calibration may be low. is there.

Therefore, the PET apparatus 100 according to the first embodiment performs the processing described below to specify the signal intensity of the peak in the histogram based only on the pair annihilation gamma ray among the pair annihilation gamma ray and the scattered gamma ray. Perform energy calibration with high accuracy.

Returning to FIG. 1, the first specific function 28a is a peak in a histogram showing the relationship between the signal intensity of the electric signal output from the photomultiplier tube 142 and the number of incident gamma rays for each of the plurality of photomultiplier tubes 142. The first signal strength, which is the signal strength of the above, is specified. The incident gamma ray refers to, for example, a gamma ray incident on the scintillator 141. Incident gamma rays include pair annihilation gamma rays and scattered gamma rays.

For example, the first specific function 28a first, for each photomultiplier tube 142, the photomultiplier tube 142 is based on the electric signals output from each of the plurality of photomultiplier tubes 142 in the clinical imaging of the subject P. The signal intensity of the electric signal output from is calculated for each incident gamma ray. Then, the first specific function 28a generates a histogram showing the relationship between the signal intensity and the number of incident gamma rays for each photomultiplier tube 142 by using the calculated signal intensity for each incident gamma ray. For example, the first specific function 28a produces a histogram 32 for a photomultiplier tube 142 as shown in the previous example of FIG.

Then, the first specific function 28a calculates the differential coefficient at each point in the histogram while moving the points on the histogram one after another from the high intensity side to the low intensity side of the signal intensity for each histogram. FIG. 7 is a diagram for explaining an example of processing executed by the first specific function 28a according to the first embodiment. For example, when the first specific function 28a generates a histogram 32 as shown in the example of FIG. 6 for a certain photomultiplier tube 142, the point is a point in the direction indicated by the arrow A as shown in the example of FIG. The differential coefficient at each point in the histogram 32 is calculated while moving. Then, the first specific function 28a identifies the point when the differential coefficient becomes 0 for the first time among the differential coefficients calculated in order from the high intensity side as the peak of the histogram based on the pair annihilation gamma ray and the scattered gamma ray. To do. That is, the first specific function 28a identifies as a peak the point where the differential coefficient first becomes 0 from the high intensity side to the low intensity side in the histogram based on the pair annihilation gamma ray and the scattered gamma ray. For example, as shown in the example of FIG. 7, the first specific function 28a specifies the point Pf when the differential coefficient becomes 0 for the first time as the peak Pf of the histogram 32.

The first specific function 28a calculates the differential coefficient while moving the points on the histogram from the high-intensity side to the low-intensity side within a range in which the number of events is larger than a predetermined threshold value. May be good. As a result, in the portion of the histogram where the first specific function 28a is affected by noise such that the number of events is less than a predetermined threshold value, the shape is deformed due to the influence of noise, and the differential coefficient becomes 0. It is possible to prevent accidentally identifying a closed point as a peak.

Then, the first specific function 28a specifies the signal strength of the specified peak as the first signal strength. For example, as shown in the example of FIG. 7, the first specific feature 28a identifies the signal strength S ₁ of a peak Pf as a first signal strength S _1.

The second specific function 28b is a pair annihilation of gamma rays incident without scattering in the histogram based on the portion of the histogram on the higher intensity side than the first signal intensity for each of the plurality of photoelectron multiplier tubes 142. The second signal strength, which is the signal strength corresponding to the energy of the gamma ray, is specified.

For example, the second specific function 28b first specifies a plurality of relationships between the signal intensity and the number of events in the portion of the histogram on the higher intensity side than the first signal intensity for each of the plurality of photomultiplier tubes 142.

FIG. 8 is a diagram for explaining an example of processing executed by the second specific function 28b according to the first embodiment. For example, as shown in the example of FIG. 8, the second specific function 28b is for a photomultiplier tube 142, signal strength and number of events in the portion of the histogram 32 of the high intensity side than the first signal strength S ₁ as the relationship between the, n-number of _(x1, y _1), to identify the _{(x 2, y 2) ····} (x n, y n). Here, x ₁ , x ₂ , ... X _n are signal strengths, respectively. Further, y ₁ , y ₂ , ... y _n are the number of events corresponding to _{x 1} , x ₂ , ... x _{n, respectively.} In the following description, when it is not necessary to distinguish each signal strength, "x" is used as the signal strength. Similarly, when it is not necessary to distinguish the number of events one by one, "y" is used as the number of events.

Then, the second specific function 28b specifies the signal intensity of the peak in the histogram based only on the pair annihilation gamma ray among the pair annihilation gamma ray and the scattered gamma ray by using the relationship between the specified plurality of signal intensities and the number of events. Here, a histogram based only on paired annihilation gamma rays is considered to approximate a Gaussian curve having a half width that is inversely proportional to the size of the square root of the number of scintillation photons. Therefore, the second specific function 28b uses n (x ₁ , y ₁ ), (x ₂ , y ₂ ) ... (x _n , y _n ) to perform curve fitting (curve fitting). By calculating an approximate curve that approximates the Gaussian curve, the signal strength of the peak in the histogram based only on the pair annihilation gamma ray is specified.

An example of a method for specifying the signal strength of such a peak will be described. For example, the Gaussian curve is represented by the following equation (1).

In equation (1), A represents the amplitude of the Gaussian function, σ represents the standard deviation, and x ₀ represents the center of the Gaussian function. Then, when the equation (1) is modified, the following equation (2) is obtained.

Then, in the second specific function 28b, n (x ₁ , y ₁ ), (x ₂ , y ₂ ) ... (x _n , y _n ) are applied to the equation (2) to fit the fitting parameters. a is, by calculating the σ and x _0, calculates a curve that approximates to a Gaussian curve. The second specific feature 28b is a x ₀ calculated, identified as signal intensity of the peak in the histogram based on only pair annihilation gamma-rays. That is, the second specific function 28b _{calculates x 0} as the second signal strength.

Here, the second specific function 28b does not use the portion of the histogram on the lower intensity side than the first signal intensity when specifying the signal intensity of the peak in the histogram based only on the pair annihilation gamma rays. The part of the histogram on the higher intensity side than the signal intensity of is used. Here, as described above, the shape of the histogram portion on the lower intensity side than the first signal intensity changes due to the influence of the scattered gamma rays, but the histogram portion on the high intensity side is the scattered gamma rays. It is considered that it is not affected by. Therefore, the second specific function 28b can accurately specify the signal intensity of the peak corresponding to the energy value (511 keV) of the pair annihilation gamma ray by using only the portion not affected by the scattered gamma ray. it can.

The second specific function 28b may add the influence of pile-up when specifying the signal strength of the peak. Here, an example of pile-up will be described. For example, the photomultiplier tube 142 receives the scintillation photon output from the scintillator 141, converts the received scintillation photon into an electric signal and outputs the photon, but the photomultiplier tube 142 is output from the scintillator 141 while the electric signal is not attenuated. When the next scintillation photon is received, the scintillation photon is converted into an electric signal, and the converted electric signal is added to the electric signal derived from the first received scintillation photon and output. That is, the photomultiplier tube 142 outputs one electric signal having a strong signal strength, which is the sum of the electric signals derived from each of the plurality of scintillation photons. Here, one electric signal having such a strong signal strength is an electric signal corresponding to one gamma ray. Such a phenomenon in which only an electric signal corresponding to one gamma ray is output from the detector module 14 even though a plurality of gamma rays are incident on the detector module 14 is called pile-up. For example, the shorter the time interval of gamma rays incident on the detector module 14, the higher the probability that pile-up will occur. FIG. 9 is a diagram for explaining a case where the second specific function 28b according to the first embodiment takes into account the influence of pile-up when specifying the signal strength of the peak. In the example of FIG. 9, a histogram 50 based on pair annihilation gamma rays and scattered gamma rays and a histogram 51 based only on pair annihilation gamma rays are shown, where the horizontal axis is the energy corresponding to the signal intensity and the vertical axis is the number of events. As described above, the portion 52 surrounded by the alternate long and short dash line on the lower intensity side than the peak of the histogram 50 based on the electrical signal obtained by the clinical imaging of the subject P may be affected by the scattered gamma rays. Further, as shown in the example of FIG. 9, the portion 53 surrounded by the dotted line on the high intensity side of the peak of the histogram 50 may be affected by the pile-up. In the portion 53, the signal strength becomes stronger due to the pile-up. Therefore, as compared with the histogram 51, the portion 53 of the histogram 50 is deformed to the side where the signal strength becomes stronger. The frequency of this pile-up is determined according to the number of incident gamma rays (count rate) per unit time. Then, if only the count rate is known, it can be estimated how much the signal strength (or the energy corresponding to the signal strength) is larger in the portion 53 than in the histogram 51.

Therefore, the second specific function 28b calculates the count rate, and based on the calculated count rate, in the portion 53, the signal strength (or the energy corresponding to the signal strength) is larger than that of the histogram 51. I guess. Then, the second specific function 28b modifies the histogram 50 so that the signal strength (or energy) is reduced by the estimated signal strength (or energy) in the portion 53. Then, the second specific function 28b specifies the second signal strength by the method as described above using the corrected histogram 50. That is, the second specific function 28b corrects the histogram 50 to remove the influence of pile-up based on the calculated count rate, and specifies the second signal strength based on the corrected histogram 50. In this way, the second specific function 28b specifies the second signal strength in consideration of pile-up. Thereby, the second signal strength can be specified more accurately.

The correction function 28c performs the above-mentioned energy calibration, and based on the second signal intensity specified for each of the plurality of photomultiplier tubes 142, the electric signal output from each of the plurality of photomultiplier tubes 142. Correct the signal strength of.

For example, the correction function 28c targets the signal strength of the electric signal that can be output from all the amplifier circuits 144 based on the width of the amplification factor of the amplifier circuit 144 and the signal strength of the electric signal output from the photoelectron multiplier tube 142. Calculated as signal strength. For example, the correction function 28c calculates the average value of the signal strengths of the electric signals output from the photomultiplier tube 142 or the amplifier circuit 144 as the target signal strength.

Then, the correction function 28c increases or decreases the amplification factor by how much from the current amplification factor because the signal strength of the electric signal output from the amplification circuit 144 becomes the target signal strength for each amplifier circuit 144. The amount of increase or decrease in the amplification factor is calculated. Then, the correction function 28c transmits the calculated increase amount or decrease amount to the amplifier circuit 144 for each amplifier circuit 144. As a result, the amplifier circuit 144 increases the amplification factor by the amount of increase when the increase amount is received, and decreases the amplification factor by the amount of decrease when the amount of decrease is received. As a result, the signal strength of the electric signals output from all the amplifier circuits 144 becomes substantially the target signal strength. In this way, the correction function 28c performs energy calibration.

10 and 11 are diagrams for explaining an example of the processing executed by the correction function 28c according to the first embodiment. For example, as shown in the example of FIG. 10, the correction function 28c calculates the target signal strength 60. Then, the correction function 28c increases the amplification factor by how much from the current amplification factor because the second signal strength of the peak 61a in the histogram 61 for a certain photomultiplier tube 142 becomes the target signal strength 60. The amount of increase in the amplification factor is calculated. Then, the correction function 28c transmits the calculated increase amount to the corresponding amplifier circuit 144.

Further, for example, as shown in the example of FIG. 11, the correction function 28c is currently used because the second signal strength of the peak 62a in the histogram 62 for the other photomultiplier tube 142 becomes the target signal strength 60. The amount of decrease in the amplification factor is calculated as to how much the amplification factor should be reduced from the amplification factor. Then, the correction function 28c transmits the calculated reduction amount to the corresponding amplifier circuit 144.

Next, an example of the flow of processing executed by the PET apparatus 100 according to the first embodiment will be described. FIG. 12 is a flowchart showing an example of a flow of processing executed by the PET apparatus 100 according to the first embodiment.

As shown in the example of FIG. 12, the system control circuit 27 determines whether or not the user has received an instruction (imaging start instruction) to start imaging in the examination of the subject P, which is input by operating the input circuit 21. Determine (step S101). If the imaging start instruction has not been received (step S101: No), the system control circuit 27 again determines in step S101 whether or not the imaging start instruction has been received.

On the other hand, when the image pickup start instruction is received (step S101: Yes), the system control circuit 27 controls the gantry device 10 so as to start the image pickup (step S102). For example, in step S102, the system control circuit 27 controls the detector module 14 to start outputting an electrical signal based on the detected incident gamma rays. Further, in step S102, the counting information collecting circuit 15 is controlled so as to generate counting information from the output signal of the detector module 14 and start storing the generated counting information in the data storage circuit 24. The imaging started in step S102 is performed until it is completed in step S106, which will be described later.

Then, the first specific function 28a of the correction circuit 28 calculates the signal strength of the photomultiplier tube 142 to which the electric signal is output for each event (step S103). The first specific function 28a includes an inspection ID which is the ID of the current inspection, an inspection date and time, an ID of the photomultiplier tube 142 which outputs the electric signal whose signal strength is calculated in step S103, and each event. The histogram generation data associated with the signal strength of the above is stored in the data storage circuit 24 (step S104).

Then, the system control circuit 27 determines whether or not the user has received the instruction to end the imaging (imaging end instruction) input by operating the input circuit 21 (step S105). When the imaging end instruction has not been received (step S105: No), the system control circuit 27 returns to step S103, and again, with respect to the photomultiplier tube 142 to which the electric signal is output, the signal strength of this electric signal. Is calculated for each event. That is, during the imaging of the subject P, the signal strength for each event is calculated in steps S103 and S104, and histogram generation data is generated.

On the other hand, when the image pickup end instruction is received (step S105: Yes), the system control circuit 27 controls the gantry device 10 so as to end the image pickup (step S106). For example, in step S106, the system control circuit 27 controls the detector module 14 to stop outputting an electrical signal based on the detected incident gamma rays. Further, in step S106, the counting information collecting circuit 15 is controlled so as to generate counting information from the output signal of the detector module 14 and stop storing the generated counting information in the data storage circuit 24.

Then, the first specific function 28a acquires the histogram generation data in which the inspection date and time is within a predetermined period from the current date and time to a predetermined time before among the histogram generation data stored in the data storage circuit 24. (Step S107).

Then, the first specific function 28a generates a histogram for each photomultiplier tube 142 using the acquired histogram generation data, and specifies the above-mentioned first signal strength for each photomultiplier tube 142. (Step S108). Here, as described above, the gain rate of the photomultiplier tube 142 changes with time. Therefore, the accuracy of the histogram generation data generated in the past inspection may be lower than that before the predetermined time from the current date and time. Therefore, the first specific function 28a generates a histogram without using the histogram generation data which may have low accuracy. As a result, it is possible to suppress a decrease in the accuracy of the generated histogram. As a result, it is possible to suppress a decrease in the accuracy of energy calibration.

In step S107, the histogram generation data may be acquired as much as the total number of events becomes equal to or greater than a predetermined threshold value. As a result, in step S108, the first specific function 28a generates a histogram using the histogram generation data for the number of events that can accurately calculate the histogram, so that an accurate histogram can be generated. Can be done.

Then, the second specific function 28b specifies the above-mentioned second signal strength for each photomultiplier tube 142 (step S109). Then, the correction function 28c executes the above-mentioned energy calibration (step S110), and ends the process.

The PET apparatus 100 according to the first embodiment has been described above. The PET apparatus 100 specifies the second signal strength of the peak in the histogram based only on the pair annihilation gamma ray among the pair annihilation gamma ray and the scattered gamma ray, and uses the specified second signal strength at the time of energy calibration. Therefore, according to the PET device 100, energy calibration can be performed with high accuracy. That is, according to the PET device 100, the intensities of the electric signals output from the plurality of photomultiplier tubes 142 can be made uniform with high accuracy.

Further, the PET device 100 executes energy calibration every time the imaging is performed. Therefore, the PET device 100 does not need to perform energy calibration for about one day during the maintenance period. Therefore, according to the PET device 100, it is possible to suppress a decrease in the operating rate.

From the above, according to the PET apparatus 100, it is possible to perform energy calibration with high accuracy while suppressing a decrease in the operating rate.

Further, as described above, the PET device 100 performs energy calibration every time the imaging is performed. Therefore, according to the PET device 100, it is possible to prevent the performance from deteriorating for a long period of time.

Further, the PET device 100 performs energy calibration without using a sealed point-shaped radiation source, a sealed linear radiation source, or the like. Therefore, according to the PET device 100, it is possible to suppress the user from feeling the complexity of having the user purchase and store a sealed point-shaped radiation source, a sealed linear radiation source, or the like.

(Second Embodiment)
In the first embodiment, the case where the PET device 100 executes the energy calibration after the imaging is completed has been described. However, the PET device 100 may perform energy calibration during imaging. Therefore, such an embodiment will be described as a second embodiment. In the second embodiment, the same processing as in the first embodiment may be designated by the same reference numerals and the description thereof may be omitted.

An example of the flow of processing executed by the PET apparatus 100 according to the second embodiment will be described. FIG. 13 is a flowchart showing an example of a flow of processing executed by the PET apparatus 100 according to the second embodiment.

The process according to the second embodiment shown in FIG. 13 is different from the process according to the first embodiment shown in FIG. 12 in that the processes of steps S107 to S110 are performed between steps S104 and S105. different.

In the second embodiment, the first specific function 28a generates a histogram in step S108 each time an electric signal is output by the photomultiplier tube 142 during imaging to specify the first signal strength. To do. Further, the second specific function 28b specifies the second signal strength in step S109 each time the first signal strength is specified by the first specific function 28a. Further, the correction function 28c performs energy calibration every time the second signal strength is specified by the second specific function 28b in step S110, and the electricity output from each of the plurality of photomultiplier tubes 142 is performed. Correct the signal strength of the signal. In this way, the PET apparatus 100 according to the second embodiment performs energy calibration in real time during imaging. Therefore, the accuracy of the counting information obtained by imaging can be further improved. Therefore, the image quality of the reconstructed PET image can be improved.

(Third Embodiment)
In the first embodiment and the second embodiment, the case where the PET device is used as the medical diagnostic imaging device has been described. However, the contents described in the first embodiment and the second embodiment can also be applied to the SPECT apparatus. This is because the SPECT device knows the energy of gamma rays depending on the nuclide used and emits X-rays of monochromatic light. Therefore, in the third embodiment, a SPECT device will be described as an example as a medical image diagnostic device to which the contents described in the first embodiment and the second embodiment are applied. In the third embodiment, the same components as those of the first embodiment and the second embodiment may be designated by the same reference numerals to omit the description or simplify the description. ..

FIG. 14 is a diagram showing an example of the configuration of the SPECT apparatus according to the third embodiment. The SPECT device according to the third embodiment includes a gantry device 10 and a console device 20.

The gantry device 10 is a device that detects gamma rays emitted from a radiopharmaceutical drug that is administered to the subject P and selectively incorporated into the biological tissue of the subject P, and collects projection data. The gantry device 10 includes a top plate 11, a sleeper 12, a sleeper drive circuit 13, a gamma camera 74, and a camera drive circuit 75. As shown in FIG. 14, the gantry device 10 has a cavity that serves as an image pickup port.

The gamma camera 74 two-dimensionally detects the intensity distribution of gamma rays emitted from the nuclide (RI: Radio Isotope) of the radiopharmaceutical drug selectively taken into the biological tissue of the subject P, and the detected two-dimensional gamma ray intensity. It is an apparatus that generates projection data by, for example, amplification processing and A / D conversion processing of distribution data. The gamma camera 74 transmits the generated projection data to the data acquisition circuit 29, which will be described later.

The camera drive circuit 75 is a device that moves the gamma camera 74 under the control of the camera control circuit 70, which will be described later. For example, the camera drive circuit 75 drives the gamma camera 74 along the inside of the image pickup port of the gantry device 10. As a result, the gamma camera 74 rotates around the subject P to generate projection data in the direction of 360 degrees.

The console device 20 accepts the operation of the SPECT device by the user, and is a nuclear medicine image (SPECT image) which is a tomographic image in which the internal distribution of the radiopharmaceutical drug administered to the subject P is drawn from the projection data collected by the gantry device 10. ) Is a device for reconstructing.

As shown in the example of FIG. 14, the console device 20 includes an input circuit 21, a display 22, a sleeper control circuit 23, a data storage circuit 24, an image reconstruction circuit 26, a system control circuit 27, and a correction circuit. It has 28, a data acquisition circuit 29, and a camera control circuit 70. Each circuit included in the console device 20 is connected via a bus.

The display 22 displays SPECT images and the like under the control of the system control circuit 27, and displays a GUI for receiving various instructions and various settings from the user via the input circuit 21.

The data collection circuit 29 collects the projection data transmitted from the gamma camera 74, and performs correction processing such as logarithmic conversion processing, offset correction, and sensitivity correction on each of the collected projection data to obtain the corrected projection data. Generate. Then, the data acquisition circuit 29 stores the generated corrected projection data in the data storage circuit 24.

The image reconstruction circuit 26 reads the corrected projection data from the data storage circuit 24, and back-projects the read corrected projection data (for example, the corrected projection data in the 360-degree direction) to create a SPECT image. To reconstruct. Then, the image reconstruction circuit 26 stores the reconstructed SPECT image in the data storage circuit 24.

The system control circuit 27 controls the operation of the gantry device 10 and the console device 20 to control the entire SPECT device. Specifically, the system control circuit 27 controls the sleeper control circuit 23 and the camera control circuit 70 to execute the projection data collection process in the gantry device 10. Further, the system control circuit 27 controls the entire image processing in the console device 20 by controlling the correction processing of the data acquisition circuit 29 and the reconstruction processing of the image reconstruction circuit 26. Further, the system control circuit 27 controls the SPECT image stored in the data storage circuit 24 so as to be displayed on the display 22.

Here, the gamma camera 74 includes the detector module 14 described in the first embodiment and the second embodiment. The scintillator 141 included in the detector module 14 according to the third embodiment converts gamma rays emitted from the internal tissue of the subject P into scintillation photons having a peak in the ultraviolet region. The photomultiplier tube 142 included in the detector module 14 converts the scintillation photons input from the scintillator 141 into an electric signal at a predetermined gain rate, and transmits the electric signal to the correction circuit 28 according to the third embodiment. To do. That is, the detector module 14 according to the third embodiment detects the gamma rays emitted from the subject P and outputs an electric signal based on the detected gamma rays.

The gamma camera 74 generates projection data from two-dimensional gamma ray intensity distribution data based on an electric signal. Then, the gamma camera 74 transmits the generated projection data to the data acquisition circuit 29.

The correction circuit 28 according to the third embodiment performs the same processing as the correction circuit 28 according to the first embodiment by using the electric signal transmitted from the gamma camera 74.

The SPECT apparatus according to the third embodiment has been described above. As described above, the correction circuit 28 according to the third embodiment performs the same processing as the correction circuit 28 according to the first embodiment. According to the SPECT apparatus according to the third embodiment, the same effect as that of the first embodiment is obtained.

In the first embodiment and the second embodiment, the PET device is taken as an example of the medical image diagnostic device, and in the third embodiment, the SPECT device is taken as an example of the medical image diagnostic device. However, the contents described in the first embodiment, the second embodiment, or the third embodiment may be applied to an X-ray CT (Computed Tomography) device as a medical image diagnostic device. For example, the contents described in the first embodiment, the second embodiment, or the third embodiment may be applied to an X-ray CT apparatus including an X-ray source that emits X-rays of monochromatic light. Further, even in an X-ray CT apparatus including an X-ray source that emits multi-colored X-rays, by using a peak corresponding to the characteristic X-ray, the first embodiment, the second embodiment, or the second embodiment. The contents described in the third embodiment may be applied. Further, the contents described in the first embodiment, the second embodiment or the third embodiment may be applied to the photon counting type X-ray CT apparatus (photon counting type X-ray CT apparatus). The plurality of detector modules included in such an X-ray CT apparatus include a radiation that has passed through a subject and is incident without scattering, and a scintillator that converts the scattered radiation of the radiation into scintillation photons. That is, the plurality of detector modules output an electric signal based on the radiation transmitted through the subject and incident without scattering and the scattered radiation of the radiation transmitted through the subject.

Further, in the first to third embodiments, the case where the photomultiplier tube is used has been described, but instead of the photomultiplier tube, an avalanche photodiode (APD) that is not affected by the magnetic field is used as a semiconductor. A silicon photomultiplier (SiPM) used as an element array may be used.

Further, in the first to third embodiments, the case where the correction circuit 28 is provided in the console device 20 has been described. However, the correction circuit 28 may be provided in the gantry device 10. Further, the counting information collecting circuit 15 may have the same function as that of the correction circuit 28 without providing the correction circuit 28.

Further, the content described in the first to third embodiments is that even when the scintillation photon from one scintillator 141 is incident on one photomultiplier tube 142, the scintillation photon from one scintillator 141 Can be applied even when is dispersed and incident on a plurality of photomultiplier tubes 142.

Further, in the first to third embodiments, the first specific function 28a determines the signal intensity of the electric signal output from the photomultiplier tube 142 and the number of incident gamma rays for each of the plurality of photomultiplier tubes 142. The case where a histogram showing the relationship between the above is generated and the first signal strength, which is the signal strength of the peak in the generated histogram, is specified has been described. However, the first specific function 28a uses the relationship between the signal intensity of the electric signal output from the photomultiplier tube 142 and the number of incident gamma rays without generating such a histogram, and uses the first signal. The strength may be specified. That is, the first specific function 28a corresponds to the peak of the number of incident gamma rays based on the relationship between the signal intensity of the electric signal output from the photoelectron multiplier tube 142 and the number of incident gamma rays without generating a histogram. The first signal strength, which is the signal strength to be used, may be specified.

Further, in the first to third embodiments, the case where the detector module 14 is an indirect conversion type detector has been described, but it may be a direct conversion type detector. For example, the detector module 14 is a direct conversion type detector composed of cadmium telluride (CdTe) -based semiconductor elements arranged in two dimensions (for example, a cadmium telluride (CdZnTe) semiconductor element). You may. The semiconductor element directly converts the incident pair annihilation gamma ray into an electric signal and outputs it. The detector module 14 composed of such a semiconductor element outputs one pulse of an electric signal (analog signal) each time a pair of annihilation gamma rays is incident.

Therefore, the detector module 14 outputs an electric signal based on the pair annihilation gamma ray emitted from the subject P regardless of whether it is an indirect conversion type detector or a direct conversion type detector.

According to the PET apparatus 100 and the SPECT apparatus of at least one embodiment described above, energy calibration can be performed with high accuracy while suppressing a decrease in the operating rate.

Although some embodiments of the present invention 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, replacements, and changes can be made without departing from the gist of the invention. These embodiments and modifications thereof are included in the scope and gist of the invention, as well as in the scope of the invention described in the claims and the equivalent scope thereof.

14 Detector module 28a First specific function 28b Second specific function 28c Correction function 100 PET device 142 Photomultiplier tube

Claims (14)

Multiple converters that output electrical signals based on incident radiation,
For each of the plurality of converters, the first signal strength, which is the signal strength corresponding to the peak of the number of radiations, is based on the relationship between the signal strength of the electric signal output from the converter and the number of incident radiations. The first specific part that identifies
For each of the plurality of converters, the signal intensity corresponding to the energy of the incident radiation without scattering is based on the relationship between the signal intensity and the number of radiations at a intensity higher than the first signal intensity. Therefore, in relation to the radiation incident on the converter without scattering and the intensity of the electric signal output from the converter due to the incident of the radiation, the signal intensity corresponds to the peak of the number of the radiations. a second specifying unit configured to specify a second signal strength Ru Oh,
A correction unit that corrects the signal strength of the electric signal output from each of the plurality of converters so that the second signal strength specified for each of the plurality of converters matches the target signal strength.
A medical diagnostic imaging device. The first specific unit identifies the first signal intensity, which is the signal intensity of the peak in the histogram showing the relationship between the signal intensity and the number of radiations, for each of the plurality of converters.
The second specific part is based on the portion of the histogram on the higher intensity side than the first signal intensity for each of the plurality of converters, and the energy of the radiation incident on the histogram without scattering is applied. Identifying the second signal strength, which is the corresponding signal strength.
The medical diagnostic imaging apparatus according to claim 1. The second specific portion is subjected to curve fitting using the relationship between the signal intensity indicated by the histogram portion on the higher intensity side than the first signal intensity and the number of incident radiations, thereby performing the second specific portion. Identify the signal strength of
The medical diagnostic imaging apparatus according to claim 2. The second specific part calculates the number of incident radiations per unit time, and based on the calculated number of incident radiations per unit time, corrects the histogram by removing the influence of pile-up. Identifying the second signal strength based on a later histogram,
The medical diagnostic imaging apparatus according to claim 2 or 3. The second specific part is formed by fitting the relationship between the signal intensity indicated by the part of the histogram on the higher intensity side than the first signal intensity and the number of incident radiations into a Gaussian curve. Identify the signal strength of
The medical diagnostic imaging apparatus according to any one of claims 2 to 4. The first specific unit identifies the first signal strength by using the relationship obtained from the electric signals output from each of the plurality of converters during the imaging of the subject.
The medical diagnostic imaging apparatus according to any one of claims 1 to 5. The first identification unit identifies the first signal strength each time the converter outputs the electrical signal.
The second specifying unit identifies the second signal strength each time the first signal strength is specified.
The correction unit corrects the signal strength of the electric signal output from each of the plurality of converters each time the second signal strength is specified.
The medical diagnostic imaging apparatus according to any one of claims 1 to 6. The first specific part uses the relationship based on the signal intensity of each incident radiation of the electric signal output from each of the plurality of converters in the inspection performed within a predetermined period, and uses the first signal. Identify the strength,
The medical diagnostic imaging apparatus according to any one of claims 1 to 7. The converter outputs the non-scattered incident radiation emitted from a radiation source with photon scattering and the electrical signal based on the scattered radiation of radiation emitted from the radiation source.
The medical diagnostic imaging apparatus according to any one of claims 1 to 8. The converter emitted radiation emitted from the subject and incident without scattering, and scattered rays of radiation emitted from the subject, or transmitted through the subject and incident without scattering. Outputs the radiation and the electrical signal based on the scattered radiation of the radiation that has passed through the subject.
The medical diagnostic imaging apparatus according to any one of claims 1 to 8. The first specific part is any one of claims 2 to 5, which specifies as the peak the point at which the differential coefficient first becomes 0 from the high intensity side to the low intensity side in the histogram. The medical diagnostic imaging apparatus described in 1. The correction unit adjusts the amplification factor of the amplifier circuit connected to the subsequent stage of each of the plurality of converters based on the second signal strength specified for each of the plurality of converters.
The medical diagnostic imaging apparatus according to any one of claims 1 to 11. The transducer outputs the electrical signal based on a pair of radiation emitted when a positron emitted in a subject combines with an electron and annihilates.
The medical diagnostic imaging apparatus according to any one of claims 1 to 8. The transducer outputs the electrical signal based on the radiation emitted from the subject.
The medical diagnostic imaging apparatus according to any one of claims 1 to 8.

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