JP2007155360A - Nuclear medical diagnosis device, and radiation detection method in nuclear medical diagnosis device - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide an inexpensive nuclear medical diagnosis device having an excellent detection characteristic of a radiation. <P>SOLUTION: In processing a detection signal, when a γ-ray 41 having incident radiation energy E0 enters the periphery 64 of the surface of a semiconductor radiation detection element 211a using a single crystal of a bromine enriched material, it is determined whether the detection signal having energy (E0-EX) smaller as much as characteristic X-ray energy EX by a characteristic X-ray 65 radiated to an adjacent semiconductor radiation detection element 211b exists in the semiconductor radiation detection element 211a or not. When the detection signal having the energy (E0-EX) exists, it is determined whether a characteristic X-ray detected simultaneously by the adjacent semiconductor radiation detection element 211b exists or not. <P>COPYRIGHT: (C)2007,JPO&INPIT

Description

  The present invention relates to a nuclear medicine diagnostic apparatus that performs imaging based on detected radiation, and a radiation detection method in the nuclear medicine diagnostic apparatus.

  Conventionally, in a nuclear medicine diagnostic apparatus, a semiconductor radiation detector for detecting radiation such as γ rays emitted from a radioactive substance administered to a subject is used. As an element of such a semiconductor radiation detector, use of thallium bromide (TlBr) as described in Patent Document 1 has been considered.

However, it has been found that a semiconductor radiation detection element made of a single crystal of high-purity thallium bromide has extremely poor energy resolution when radiation such as γ rays is incident. For example, when cesium 137 ( ¹³⁷ Cs) is used as the radiation source, the spectrum of the peak value of the pulse of the detection signal becomes broad, the first peak at the 662 keV position of the single energy to be detected, and the first peak The second peak at 589 keV adjacent to the peak does not appear clearly. That is, it has been found that a semiconductor radiation detection element made of a single crystal of high-purity thallium bromide has a problem that it has a poor energy resolution and cannot be put into practical use in terms of detection accuracy.

In order to solve this problem, a semiconductor radiation detector having a structure in which both surfaces of a single crystal of thallium bromide containing 0.1 to 20.0% of trivalent thallium are sandwiched between a first electrode and a second electrode Was realized. As a result, the resistivity of the single crystal of thallium bromide is improved (that is, the resistance value per unit volume of the single crystal is reduced), so that a high voltage can be applied to the semiconductor radiation detector, and Single crystal lattice defects are also reduced. And since the energy resolution and sensitivity of a semiconductor radiation detector can be improved, it became possible to improve the detection accuracy of a semiconductor radiation detector as a result (refer patent document 1).
Japanese Patent Laying-Open No. 2005-223209 (paragraphs 0025 to 0030)

However, in such a semiconductor radiation detector using thallium bromide, the radiation source is not limited to cesium 137, and a single energy position to be detected with respect to incidence of all single energy radiations. And a second peak with an energy of 73 keV to 83 keV lower than the first peak appear adjacent to each other. Such a phenomenon significantly deteriorates the imaging performance of the nuclear medicine diagnostic apparatus.
That is, as a result of the pulse height analysis of the detection signal, when the second peak appears on the lower side adjacent to the energy position of the first peak, both peaks become the original single energy radiation detection signal, As a result, the scattered components are also counted, and the captured image is unclear and the inspection accuracy is lowered. If only the detection signal of the first peak position is the original single energy radiation detection signal, the imaging sensitivity is reduced, and as a result, it takes time to obtain a clear captured image.

  The present invention has been made in view of the above problems, and an object of the present invention is to provide a nuclear medicine diagnostic apparatus having excellent radiation detection characteristics and good imaging performance.

  In order to solve the above-mentioned problem, in the present invention, radiation emitted from the body of a subject or radiation transmitted through the body of the subject is detected using a plurality of semiconductor radiation detectors, and a detection signal of the radiation is detected. A nuclear medicine diagnostic apparatus for obtaining a captured image of the subject on the basis of an incident radiation of energy (E0) incident on a first semiconductor radiation detector among the plurality of semiconductor radiation detectors The detected radiation energy (E1) is detected at a second semiconductor radiation detector adjacent to the first semiconductor radiation detector by generating the incident radiation energy (E0) and the incident radiation. Correction means for determining whether or not the difference from the characteristic X-ray energy (EX) is equal and determining whether or not to correct the detection signal of the first semiconductor radiation detector based on the determination result. Preparation It was the thing.

The correction means includes, for example, a detection radiation energy (E1) detected by the first semiconductor radiation detector among the arrayed semiconductor radiation detectors arranged in two dimensions or three dimensions, and the adjacent second semiconductor radiation. If the sum of the predetermined energy characteristic X-ray energy (EX) detected by the detector at approximately the same time coincides with the incident radiation energy (E0), the first
When the incident radiation of energy (E0) can be corrected and counted at the signal detection location of the semiconductor radiation detector of FIG. 5, and the incident radiation generates characteristic X-rays and exhibits a lower detected radiation energy (E1) Can also be determined as detection of regular incident radiation.

  According to the present invention, even in the case of a semiconductor radiation detector in which incident radiation generates characteristic X-rays, the detection characteristics of the semiconductor radiation detector with respect to the incident radiation can be improved, and the nuclear medicine diagnosis with good imaging performance. An apparatus can be provided.

  The nuclear medicine diagnostic apparatus according to the embodiment of the present invention will be described below.

<< First Embodiment >>
A first embodiment of a nuclear medicine diagnosis apparatus according to the present invention will be described with reference to FIGS. This embodiment is a γ camera imaging system as a nuclear medicine diagnostic apparatus using a semiconductor radiation detector.
As shown in FIG. 1, the γ camera imaging system 1 detects γ rays emitted from a radiopharmaceutical administered to a subject P and obtains a captured image. For imaging, a radiopharmaceutical containing a radionuclide that emits a single energy (E0) gamma ray 41 is administered to the subject P, and the single energy released from the affected area of the subject on which the drug is concentrated. Gamma rays 41 are used.

(Configuration of γ camera imaging system)
The γ camera imaging system 1 mainly includes a camera 3, a data collection device 5a, a data analysis device 5b, and a monitor 7.
The camera 3 detects γ rays emitted from the body of the subject. The camera 3 is attached to the camera base 4 so that it can be driven in the X and Y directions shown in FIG. Detection information when γ rays are detected by the camera 3 is transmitted to the data collection device 5a.
In FIG. 1, the side where the subject P stands with respect to the camera 3 is referred to as the front, and the opposite direction is referred to as the rear.
The data collection device 5a is a device for accumulating detection information data, and processes the detection information data obtained by the camera 3 for each semiconductor radiation detector 13 (see FIGS. 3 and 4) described later, Accumulate. The data analysis device 5b analyzes the data stored in the data collection device 5a, generates a captured image, and displays it on the monitor 7.

(Camera configuration)
Next, the configuration of the camera 3 of the γ camera imaging system 1 will be described with reference to FIG.
As shown in FIG. 2, a collimator 11 made of a lead plate is provided along the imaging surface on the incident side of the γ ray 41 of the camera 3. The collimator 11 is provided with a large number of windows 11a drilled in the horizontal direction and realizes good directivity with respect to γ rays incident on a semiconductor radiation detector 13 which will be described later.

The individual semiconductor radiation detectors 13 are two-dimensionally arrayed on the detector connector board 15 along the imaging surface of the camera 3 so as to correspond to the window portion 11a behind the collimator 11, Gamma rays 41 emitted from the body of the examiner and the semiconductor radiation detector 13 (element)
The detection signal of the characteristic X-ray generated in step 1 is output.
An ASIC module board 18 on which an analog integrated circuit (hereinafter referred to as an analog ASIC) 24 is mounted is disposed behind the detector connector board 15. Detection signals of γ rays and X rays from the semiconductor radiation detector 13 are transmitted to the analog ASIC 24 via the internal wiring (not shown) of the ASIC module board 18.

The ASIC module board 18 on which the semiconductor radiation detector 13, the detector connector board 15, and the analog ASIC 24 are mounted is covered with a light shielding shield case 12 so as to be shielded from light.
The light shielding shield case 12 corresponds to the housing of the present invention.
An output connector C2 is provided on the rear surface of the light shielding shield case 12, and a signal from the output terminal of the ASIC module board 18 is sent to the data collection device 5a via the output connector C2.

(Configuration of detector connector board and ASIC module board)
Next, the detector connector board 15 and the ASIC module board 18 will be described in detail with reference to FIG.
The detector connector board 15 includes a connector C1 for connecting the semiconductor radiation detector 13, and a capacitor 22 and a resistor 23 provided corresponding to each semiconductor radiation detector 13. The detection signals from the individual semiconductor radiation detectors 13 are differentiated through the capacitors 22 and the resistors 23, and are supplied to one system of the analog detection circuit 27 described later on the analog ASIC 24 so that predetermined signal processing is performed. It has become.

The ASIC module board 18 mainly includes an analog ASIC 24, an analog / digital converter (hereinafter referred to as ADC) 25, and a digital integrated circuit (hereinafter referred to as digital ASIC) 26.
The analog ASIC 24 includes a large number of analog detection circuits 27 that process detection signals output for each semiconductor radiation detector 13.

One analog detection circuit 27 is composed of a fast system and a slow system, and one analog detection circuit 27 is provided corresponding to one set of the semiconductor radiation detector 13, the capacitor 22, and the resistor 23.
The first system is provided with a timing pick-off circuit 24a that generates a timing signal in order to determine the time when the γ-ray and the X-ray are detected. The timing signal is input directly to the digital ASIC 26 where it is signal processed.

The slow system is a band that performs band-pass processing to remove noise for the purpose of obtaining peak values of detected γ-rays and X-rays, a charge amplifier 24b as a preamplifier, a polarity amplifier 24c as a linear amplifier. A pass filter (discriminator circuit) 24d and a peak hold circuit (wave height analysis circuit) 24e for capturing the peak value are provided in this order. By the way, the slow system is named “slow” because it takes a certain amount of processing time to obtain the peak value.
The capacitor 22, the resistor 23, and the analog ASIC 24 constitute a signal processing circuit according to the present invention. In particular, the peak hold circuit 24e constitutes a pulse height analysis circuit according to the present invention.

The slow output signal has an analog peak value. The analog peak value is converted from an analog signal to a digital signal by the ADC 25 and input to the digital ASIC 26, where the signal is processed.
Note that the ADC 25 converts the analog peak value into a digital peak value of 8 bits (256 numeric strings from 0 to 255). For example, when the single γ-ray energy released from the radiopharmaceutical is 662 KeV, the corresponding peak value corresponds to the digital value 255.

The digital ASIC 26 has a 500 MHz clock, and determines the detection time of γ rays and X rays based on the input timing signal. Since the timing signal is based on the first signal of the analog ASIC 24, a time close to the true detection time can be set as the detection time.
The detection time is determined by the LET method that determines the timing at the rising point of the waveform or the CFD method that determines the timing when the waveform rises by 20%.

  By the way, regarding the first system that generates the timing signal, each semiconductor radiation detector 13 has a one-to-one correspondence with each connection terminal (input terminal) (not shown) of the digital ASIC 26. As for the slow system for sampling the crest value, each one of the semiconductor radiation detectors 13 has a one-to-one correspondence with each connection terminal (input terminal) (not shown) of the digital ASIC 26.

In this way, the digital ASIC 26 sets the peak value of the detection signal, the detection time determined by the timing signal, and the ID of the semiconductor radiation detector that uniquely identifies each semiconductor radiation detector 13. As a signal, correction processing for counting loss due to leakage characteristic X-rays, which will be described later in detail, is performed. The outline of the processing will be described below.
When the semiconductor radiation detector 13 outputs a detection signal corresponding to the incident γ-ray energy (E0) as a radiation detection signal, the semiconductor radiation detector 13 determines that the incident γ-ray of energy (E0) has been detected. To do.

Further, a certain semiconductor detector 13 emits radiation (E0-EX) with a predetermined energy lower than the incident γ-ray energy (E0), that is, energy (E0-EX) smaller than the characteristic X-ray energy (EX) due to the characteristic X-ray. It is assumed that the detected detection signal (first detection signal) is output. At the same time, the second detection signal is output from the semiconductor radiation detector 13 adjacent to the semiconductor radiation detector 13 that has output the first detection signal, and the detection signal corresponds to the characteristic X-ray energy (EX). The semiconductor radiation detector 13 that has output the first detection signal determines that incident γ-rays of energy (E0) have been detected.

When the first detection signal and the second detection signal do not meet the above conditions, the first detection signal is a scattered component, and it is not determined that an incident γ ray of energy (E0) has been detected.
The digital ASIC 26 has a function of instructing the ADC 25 about the output timing of the digitized peak value output from the ADC 25.

(Configuration of semiconductor radiation detector)
Next, the configuration of the semiconductor radiation detector 13 will be described with reference to FIG.
The semiconductor radiation detector 13 is a single crystal of bromine-enriched thallium bromide (hereinafter referred to as bromine-enriched product) obtained through the process of enriching bromine with thallium bromide described in Patent Document 1. This is a configuration in which a semiconductor radiation detection element 211 using a single crystal) and an anode electrode A and a cathode electrode C are arranged on both sides thereof.

The semiconductor radiation detection element 211 of FIG. 4 is a rectangular rod body in which a bromine-enriched single crystal is sliced along a predetermined lattice plane and has a substantially square cross section. In the present embodiment, one side of the cross section of the semiconductor radiation detection element 211 is set to about 1.0 mm. In addition, trivalent thallium is contained in the bromine-enriched single crystal. As is well known, this trivalent thallium is produced by bringing bromine into contact with primary thallium bromide. In addition to tertiary thallium bromide, hexacoordinate and tetracoordinate bromothallium (III ) It is considered to be contained in the single crystal in the form of thallium (I) salt of a complex group. Specific examples of those containing trivalent thallium include TlBr ₃ , 3TlBr · TlBr ₃ , Tl ₃ [TlBr ₆ ], TlBr · TlBr ₃ , Tl [TlBr ₄ ] and the like.

  The cathode electrode C and the anode electrode A are applied with a voltage for collecting charges for collecting charges generated by the semiconductor radiation detection element 211 by receiving γ rays, for example, by a DC power supply V of about 300V to 800V. It is like that. For the cathode electrode C and the anode electrode A, for example, a highly conductive film such as Pt (platinum), Au (gold), or In (indium) is used.

(Description of operation of γ camera imaging system)
Next, the operation of the γ camera imaging system 1 will be described with reference to FIGS. 4 to 9 (refer to FIGS. 1 and 2 as appropriate).

(Operation of semiconductor radiation detector)
First, the response of the semiconductor radiation detector 13 when γ rays or X-rays enter the semiconductor radiation detector 13 will be described.
As shown in FIG. 4, when γ-rays or X-rays are incident on the semiconductor radiation detection element 211, they interact with the γ-rays or X-rays, thereby schematically showing (+) and (−) in the figure. Generate the hole-electron pair shown. At this time, holes (+) and electrons (−) are generated in pairs in amounts proportional to the energy of γ rays or X rays. On the other hand, since a voltage for collecting charges is applied between the anode electrode A and the cathode electrode C by a DC power supply V of about 300V to 800V, the holes are attracted to the cathode electrode C and moved. The electrons are attracted to the anode electrode A and move.

However, when a single energy gamma ray is incident on the semiconductor radiation detection element 211 using a bromine-enriched single crystal, which is a feature of the present embodiment, a gamma ray having an energy of 73 keV to 83 keV lower than the incident energy is detected. Is done. This phenomenon is related to the emission of characteristic X-rays generated when γ rays are incident on thallium bromide.
In FIG. 5, the anode electrode A and the cathode electrode C are omitted, and only the semiconductor radiation detection element 211 is shown, but one semiconductor radiation detector 13 is represented.

Single energy (E0) γ rays 41 are incident on the semiconductor radiation detection element 211, and the energy (E0) is absorbed by the element inside 61 of the semiconductor radiation detection element 211. Hereinafter, this energy (E0) is referred to as incident radiation energy (E0).
At this time, a part of the incident radiation energy (E0) is re-emitted as the energy (EX) of the characteristic X-ray 62. Hereinafter, the energy (EX) of the characteristic X-ray 62 is referred to as characteristic X-ray energy (EX). As shown in FIG. 5A, the γ-ray 41 is incident on a semiconductor radiation detection element 211 to generate a characteristic X-ray 62, and the characteristic X-ray 62 is completely generated in the same semiconductor radiation detection element 211 as it is. If detected, the detection signal of the semiconductor radiation detection element 211 becomes a signal having a peak value corresponding to the incident radiation energy (E0), and there is no problem in detection accuracy.

  On the other hand, as shown in FIG. 5B, the γ-ray 41 is incident on a semiconductor radiation detection element 211a (corresponding to a first semiconductor radiation detector), and is absorbed by the surface vicinity 64 of the semiconductor radiation detection element 211a. In such a case, the generated characteristic X-ray 65 enters the adjacent semiconductor radiation detection element 211b (corresponding to the second semiconductor radiation detector). That is, in this case, the energy (E1) [detected radiation energy (E1)] indicated by the peak value of the detection signal of the semiconductor radiation detection element 211a is lower than the incident radiation energy (E0) by the characteristic X-ray energy (EX). become. That is, E1 = (E0−EX). For example, when the incident radiation energy (E0) is 662 keV and the characteristic X-ray energy (EX) is 73 keV, the detected radiation energy (E1) of the semiconductor radiation detection element 211a is 589 (= 662-73) keV.

In the case of FIG. 5B, a detection signal having a peak value corresponding to the characteristic X-ray energy (EX) by the characteristic X-ray 65 is output from the semiconductor radiation detection element 211b adjacent to the semiconductor radiation detection element 211a at the same timing. Will be. Such a phenomenon is that when a single energy gamma ray is incident on the semiconductor radiation detector 13 using a bromine-enriched single crystal, it is detected as an energy that is 73 keV to 83 keV lower than the incident radiation energy (E0). Become.
The difference between the characteristic X-ray energy (EX) of 73 keV or 83 keV depends on whether the incident γ-ray 41 ejects electrons of the K shell or L shell of thallium atoms. For example, when characteristic X-rays are generated in the K shell, the characteristic X-ray energy (EX) is 72.872 keV (about 73
kev), when characteristic X-rays are generated in the L shell, the characteristic X-ray energy (EX) is 82.576.
keV (about 83 keV).

Since the characteristic X-ray energy (EX) is as small as 100 keV or less, when the cross-sectional dimension of the semiconductor radiation detector 13 is 1 mm or more in both the X direction and the Y direction, the semiconductor radiation detector adjacent to the characteristic X-ray is used. The probability of being absorbed by the semiconductor radiation detector 13 that has passed through 13 and further away is negligible.
Incidentally, if the semiconductor radiation detection element is miniaturized in order to increase the resolution of the image, there is a high possibility that characteristic X-rays are emitted to other semiconductor radiation detection elements. In the case of such a small semiconductor radiation detection element, correction processing may be performed with the adjacent range widened.

  Then, the semiconductor radiation detector 13 outputs a detection signal indicating the magnitude of the detected radiation (γ-ray or X-ray) energy according to the amount of electrons collected at the anode electrode A, that is, the magnitude of the electric charge.

(Signal processing in analog ASIC)
Upon receiving this detection signal, the analog detection circuit 27 (see FIG. 3) amplifies the detection signal with the charge amplifier 24b and the polarity amplifier 24c, and passes the band-pass filter 24d to remove the noise to detect the detection signal from the peak hold circuit. An analog peak value signal is generated by analyzing at 24e. The analog peak value signal is output to the ADC 25 and converted into a digital peak value signal, and the digitally converted peak value signal is output to the digital ASIC 26. On the other hand, a timing signal is output to the digital ASIC 26 from the timing pick-off circuit 24a (see FIG. 3) of the analog detection circuit 27.

(Correction process)
Next, the flow of the correction process for counting loss due to the leakage characteristic X-ray performed in the digital ASIC 26 will be described with reference to FIGS.
In the γ camera imaging system of FIG. 1, γ rays of incident radiation energy (E0) are incident on the camera 3 from the inside of the subject P. The characteristic X-ray energy generated by the γ rays in the semiconductor radiation detector 13 is defined as characteristic X-ray energy (EX).
Detection information obtained along with radiation detection by the semiconductor radiation detector 13 is detected coordinates (X, Y).
), Detected radiation energy (E), and detection time (T). The detected coordinates (X, Y) are
This corresponds to the ID of each semiconductor radiation detector 13.

FIG. 8 shows the peak value distribution of the detection signal output from the slow system of the analog detection circuit 27. The horizontal axis represents γ-ray (or X-ray) energy, and the vertical axis represents the detection signal count.
The first detection signal that is a candidate for detection of incident radiation is a detection signal that exceeds a threshold Eth greater than the Compton edge 76. The incident radiation energy peak 72 is a peak corresponding to the case where the semiconductor radiation detector 13 detects the incident radiation energy (E0). The low energy component peak 73 is a peak corresponding to the case where the characteristic X-ray leaks from the semiconductor radiation detector 13. The second detection signal 75 is a peak corresponding to the case where characteristic X-rays are detected.

Region 74 shown by a dot exceeding the threshold Eth shows a scattering component, if it is determined that the detection signal of the peak value exceeding the threshold value E _th discovered all γ-rays of the incident radiation energy (E0), scattered component 74 Are also counted, which deteriorates the image.
The components indicated by broken lines including the scattering component 74 are due to scattering and the like, and are referred to as noise 71 in handling the detection signal when a captured image is generated by radiation.

In FIG. 7, assuming that the first semiconductor radiation detector (D1) 80 at coordinates (X1, Y1) has detected a radiation with energy higher than a predetermined threshold (Eth), detection information associated with the first detection signal. Are detected coordinates (X1, Y1), detected radiation energy (E1), and detected time (T1).
Note that the predetermined threshold (Eth) is an appropriate value that exceeds the value of the Compton edge 76 by checking the energy spectrum in advance when the semiconductor radiation detector 13 detects single energy γ rays from the radionuclide used. Choose the correct value.

In FIG. 7, when the second semiconductor radiation detector (D2) 81 in the area indicated by the oblique lines adjacent to the coordinates (X1, Y1) outputs the second detection signal with respect to the first detection signal. The detection information associated with the second detection signal includes detection coordinates (X2, Y2) and detection radiation energy (E2).
The detection time (T2).
In FIG. 7, any one of the eight detectors adjacent to the first semiconductor radiation detector (D1) 80 is a detector 81. When the semiconductor radiation detectors are arranged in three dimensions, the adjacent range is widened.

In step S1, first, it is checked whether or not the detected radiation energy (E1) of the first detection signal exceeds a threshold value (Eth). If it exceeds the threshold (Eth) (Yes), the process proceeds to step S2, and if it does not exceed (No), step S1 is repeated.
In step S2, it is determined whether or not the detected radiation energy (E1) matches the incident radiation energy (E0), that is, whether [E1 (D1) = E0].
Here, if the detected radiation energy (E1) is within a range that is twice the half-value width of the incident radiation energy peak 72 shown in FIG.
When the detected radiation energy (E1) of the first detection signal matches the incident radiation energy (E0), that is, when [E1 (D1) = E0] (Yes), the process proceeds to step S6. 1 semiconductor radiation detector (D1) 80 determines that regular radiation detection has been performed.

  On the other hand, if the detected radiation energy (E1) does not match the incident radiation energy (E0) in the determination in step S2 (No), that is, if [E1 (D1) ≠ E0] (No), the process proceeds to step S3. Advancing, whether or not the detected radiation energy (E1) of the first detection signal corresponds to the incident radiation energy (E0) minus the characteristic X-ray energy (EX), that is, [E1 (D1) = E0 -EX] or not. Here again, as in step S2, it is determined whether or not there is a value within a predetermined range.

When the detected radiation energy (E1) of the first detection signal is not the incident radiation energy (E0) minus the characteristic X-ray energy (EX), that is, [E1 (D1) ≠ E0
In the case of -EX] (No), the process proceeds to step S7, where the first detection signal is determined to be a scattered component, and is removed from the detection information accumulated in the data collection device 5a.

  On the other hand, in the determination of step S3, when the detected radiation energy (E1) of the first detection signal is obtained by subtracting the characteristic X-ray energy (EX) from the incident radiation energy (E0), that is, [E1 (D1 ) = E0−EX] (Yes), the process proceeds to step S4.

  In step S4, the second semiconductor radiation detection adjacent to the first semiconductor radiation detector (Dl) at the same time as the detection time (T1) of the first detection signal in the first semiconductor radiation detector (Dl). It is determined whether or not the device (D2) has obtained the second detection signal, that is, whether [T1 (D1) = T2 (D2)]. This is determined by whether the detection time (T2) of the second detection signal is within a predetermined time window (for example, 10 nanoseconds) with respect to the detection time T1. As shown in FIG. 7, the second semiconductor radiation detector (D2) adjacent to the first semiconductor radiation detector (D1) 80 is the semiconductor radiation detector 81 in the region indicated by hatching.

If the time window is widened, the cost of hardware for simultaneous measurement can be reduced, but the probability of accidental determination as simultaneous increases.
When the detection time (T2) of the detected radiation energy (E2) by the adjacent second semiconductor radiation detector (D2) 81 is not the same time at the same time as the first detection signal, that is, [T1 (D1) ≠ In the case of T2 (D2)] (No), the process proceeds to step S7 and is excluded from the detection information as a scattering component.

On the other hand, in the determination of step S4, the detection time (T1) of the first detection signal and the detection time (T2) of the second detection signal of the adjacent second semiconductor radiation detector (D2) are the same time. In the case, that is, [T1 (D1) = T2 (D2)] (Yes), the process proceeds to step S5. In step S5, it is determined whether or not the detected radiation energy (E2) of the second detection signal matches the characteristic X-ray energy (EX), that is, [E2 (D2) = EX]. Here again, as in step S2, it is determined whether or not there is a value within a predetermined range.

When the detected radiation energy (E2) of the second detection signal does not match the characteristic X-ray energy (EX), that is, when [E2 (D2) ≠ EX] (No), the process proceeds to step S7, and the first Is detected as a scattered component and removed from the detection information.
In step S5, if the detected radiation energy (E2) of the second detection signal is the same as the characteristic X-ray energy (EX), that is, if [E2 (D2) = EX] (Yes), step S6. The detection information of the first detection signal is determined to detect the normal radiation at the coordinates (X1, Yl).
After steps S6 and S7, the correction process for the series of first detection signals ends.

  6 is performed by the digital ASIC 26 (not shown) provided on the ASIC module board 18 in the camera 3 of FIG. 2 in this embodiment, the γ camera imaging system 1 of FIG. The data collecting device 5a may be used.

  FIG. 9 is a diagram schematically showing the imaging results before and after performing the count drop correction process using the leakage characteristic X-ray described in FIG. (A) shows the position of the semiconductor radiation detector which output the detection signal of the detection radiation energy exceeding a threshold value (Eth). Among them, the position indicated by dots is the position where the incident radiation energy (E0) is detected, and the position indicated by diagonal lines including the double diagonal lines is greater than the threshold (Eth) and smaller than the incident radiation energy (E0). Further, the position with double diagonal lines indicates the detection position from which the first detection signal that has undergone the process from step S3 to steps S4 and S5 to step S6 in the flowchart of FIG. 6 is output.

When correction processing for counting loss due to leakage characteristic X-rays is performed, the single-shaded portion in FIG. 9A is removed as the scattering component 74 (see FIG. 8), and the double-shaded portion is determined to be normal radiation detection. Therefore, an area as indicated by a dot in (b) is an area where it is determined that regular radiation detection has been performed.
That is, before performing the count loss correction process by the leakage characteristic X-ray, as shown in FIG. 9A, there is a portion where it is not clear whether the hatched portion is a scattering component, and the wave exceeding the threshold Eth If the high-value detection signal is used as it is, the captured image becomes unclear. On the other hand, after performing the correction process, as shown in (b), since the position of the dot is a position determined to be normal radiation detection, a clear captured image can be obtained.

Note that the digital ASIC 26 in the present embodiment constitutes the correcting means of the present invention. In particular, step S3 to step S7 in the flowchart constitute the correcting means of the present invention.
Step S1 in the flowchart corresponds to the first step of the present invention, step S3 corresponds to the second step, steps S4 and S5 correspond to the third step, and step S6 corresponds to the fourth step.

  As described above, according to the present embodiment, correction of counting omission due to leakage characteristic X-rays is performed on a detection signal having a peak value equal to or higher than the threshold Eth, and detection detection of normal incident radiation is performed, so that a clear image is obtained. be able to. Further, in FIG. 8, the imaging sensitivity can be improved only for the incident radiation energy peak 72 as compared with the case where the normal detection is determined by the energy window having a predetermined half width.

Further, in the γ camera imaging system 1 according to the first embodiment, since the semiconductor radiation detection element 211 is composed of a single crystal of bromine enriched thallium bromide, a single crystal of cadmium telluride was used. Compared to a conventional γ camera imaging system, the cost is much lower.
Since the semiconductor radiation detection element 211 used in this embodiment generates charges in response to light, the light shielding shield case 12 is made of a light shielding material such as aluminum or aluminum alloy. It is configured so as to eliminate a gap through which light enters.

Note that the counting drop correction method using the leakage characteristic X-ray in the present embodiment is not limited to the flowchart shown in FIG.
A first determination (corresponding to step S4) of whether or not adjacent semiconductor radiation detectors have detected radiation substantially simultaneously among a plurality of semiconductor radiation detectors, and then one of the adjacent semiconductor radiations is performed. A second determination (corresponding to step S5) is made as to whether or not the energy of the radiation detected by the detector is equal to the energy (EX) of the characteristic X-ray, and the two adjacent semiconductor radiation detectors detected. Whether the sum of the energy of the radiation is equal to the energy (E0) of the incident radiation, or the energy of the radiation detected by the other semiconductor radiation detector adjacent to the energy of the incident radiation (E0) and the characteristic X-ray A third determination is made as to whether or not the energy (EX) difference is equal to the difference in energy (EX) (note that the latter part of the third determination corresponds to step S3). It may be determined whether the detection of the radiation, i.e. whether it detects a normal radiation based on. Of course, the order of each determination may be changed.
The data processing apparatus according to the claims corresponds to the digital ASIC 26 in the present embodiment. Further, the radiation detection determination apparatus according to the claims corresponds to the digital ASIC 26 in the present embodiment.

<< Second Embodiment >>
Next, a second embodiment of the nuclear medicine diagnosis apparatus according to the present invention will be described with reference to FIGS.
In addition, the same code | symbol is attached | subjected about the same structure as 1st Embodiment, and description is abbreviate | omitted.
This embodiment is a PET (Positron Emission Computed Tomography) system as a nuclear medicine diagnostic apparatus using a semiconductor radiation detection element.

The subject is administered fluoro-dioxy-glucose (FDG) containing fluorine 18 ( ¹⁸ F) with a half-life of 110 minutes in advance. This FDG concentrates on the affected area of the subject. When a positron emitted from fluorine 18 ( ¹⁸ F) annihilates with an electron, a 511 KeV single energy γ-ray emitted at an angle of 180 ° to each other is detected by the PET system.

(PET system)
The PET system 30 mainly includes a camera 31, a data processing device 32, a monitor 33, and a bed 34.
The camera 31 detects γ rays emitted from the body of the subject and has a cylindrical void.
The bed 34 can be driven so as to move in the axial direction of the cylinder within the void of the camera 31 with a subject not shown.

  The data processing device 32 also serves as a data storage device, and data for each γ-ray, X-ray peak value and detection time detected by the camera 31 is for each semiconductor radiation detector 13A (see FIGS. 12 and 13) to be described later. It accumulates in. The data processing device 32 incorporates a simultaneous measuring device 32a (see FIG. 11) described later. The data processing device 32 performs simultaneous measurement every predetermined time window by the simultaneous measuring device 32a, specifies the generation position of the γ-ray, generates a captured image based on the specified position, and displays it on the monitor 33. It is like that.

  As shown in FIG. 11, a large number of semiconductor radiation detector units 16, for example, tens of thousands to hundreds of thousands are arranged in the circumferential direction on the radially outer side of the cylindrical space of the camera 31. The semiconductor radiation detector unit 16 outputs a detection signal for γ rays emitted from the body of the subject P positioned at the center of the cavity of the camera 31. A signal from the semiconductor radiation detector unit 16 is input to the simultaneous measurement device 32a.

(Configuration of semiconductor radiation detector unit)
Next, the semiconductor radiation detector unit 16 will be described in detail.
As shown in FIG. 12, the semiconductor radiation detector unit 16 mainly includes a number of semiconductor radiation detectors 13A, an analog ASIC 24, and a digital ASIC 26A.
The configuration of the semiconductor radiation detector unit 16 of this embodiment is close to the block configuration of the detector connector board 15 and the ASIC module board 18 shown in FIG. 3 of the first embodiment. The difference is that the semiconductor radiation detector 13 is replaced with a semiconductor radiation detector 13A, and the digital ASIC 26 is replaced with a digital ASIC 26A.

(Configuration of semiconductor radiation detector)
As shown in FIG. 13, the semiconductor radiation detector 13 </ b> A is a semiconductor radiation detector that uses a bromine-enriched single crystal obtained through the step of enriching bromine with the thallium bromide described in Patent Document 1. The element 211 has a configuration in which an anode electrode A and a cathode electrode C are disposed on both sides thereof, and a plurality of groups are connected in parallel.

  That is, as shown in FIG. 13, the semiconductor radiation detector 13A has a five-layer laminated structure in which the semiconductor radiation detection element 211 is sandwiched between the cathode electrode C and the anode electrode A. Each anode electrode A is connected in parallel to one anode line to one cathode line. The semiconductor radiation detector unit 16 shown in FIGS. 11 and 12 is provided with a plurality of such semiconductor radiation detectors 13A in two dimensions or three dimensions.

  A semiconductor radiation detection element 211 shown in FIG. 13 is obtained by slicing a bromine-enriched single crystal along a predetermined lattice plane, and its planar shape is a rectangular plate-like body. In the second embodiment, the thickness of the semiconductor radiation detection element 211 is set to about 1.0 to 0.5 mm. This single crystal of bromine enrichment is the same as that described in the first embodiment.

  The cathode electrode C and the anode electrode A are applied with a voltage for collecting charges for collecting charges generated by the semiconductor radiation detecting element 211 by receiving γ rays by a DC power source (not shown) of about 300 V to 800 V, for example. It has become so. For the cathode electrode C and the anode electrode A, for example, a highly conductive film such as Pt (platinum), Au (gold), or In (indium) is used.

  Incidentally, in the semiconductor radiation detector 13A in this embodiment, since the anode electrodes A and the cathode electrodes C are connected in common, each layer detects radiation independently of the other layers. Absent. In other words, when the incident gamma ray interacts with the semiconductor radiation detection element 211, whether it is caused by the uppermost semiconductor radiation detection element 211, the lowermost semiconductor radiation detection element 211, or the like. The configuration is not determined. Of course, it can also be set as the structure detected independently for every semiconductor radiation detection element 211 of each layer. The reason why the semiconductor radiation detection element 211 has such a five-layer structure is that the semiconductor radiation detection element 211 and the γ-ray are reduced by reducing the amount of γ rays that pass through while increasing the charge collection efficiency. This is to increase the interaction with. That is, this is to increase the count number of detection of γ rays.

  With the configuration of the semiconductor radiation detector 13A having a laminated structure as shown in FIG. 13, a better crest value rising speed (that is, a steep rise) and a more accurate crest value can be obtained, and semiconductor radiation detection can be performed. The number of γ rays that interact with the semiconductor material of the element 211 (that is, the count number) can also be increased. As a result, the sensitivity of the semiconductor radiation detector 13A can be increased.

(Mounting form of semiconductor radiation detector unit)
Next, a configuration for mounting the semiconductor radiation detector unit 16 on the camera 31 will be described with reference to FIG.
A state when the semiconductor radiation detector unit is attached to the camera 31 is shown in FIG. FIG. 2B is a cross-sectional view taken along the line AA in the center of the camera 31. As shown in (b), the semiconductor radiation detector unit 16 is attached to the camera 31 via a unit support member 31b. The semiconductor radiation detector unit 16 is mounted on the camera 31 with one end supported by a unit support member 31b. The unit support member 31b has a hollow disk shape (donut shape), and has a large number of windows for mounting the semiconductor radiation detector units 16 in the circumferential direction (that is, the number of the semiconductor radiation detector units 16 to be mounted). I have.

  Thus, since the camera 31 supports the semiconductor radiation detector unit 16 at one end, a flange 35a serving as a stopper is provided on the front side of the housing 35 of the semiconductor radiation detector unit 16 in the body axis direction. Incidentally, when the semiconductor radiation detector units 16 are arranged as closely as possible in the circumferential direction, the flange 35a on the side (radially inner side) along the cylindrical void of the camera 31 becomes an obstacle. Therefore, the flange 35a at this hindrance may be eliminated from the housing 35 and the radially outer flange 35a may be left.

  When the semiconductor radiation detector unit 16 is mounted on the camera 31, the lid 31a shown in FIG. 14B is removed to expose the unit support member 31b, from which the semiconductor radiation detector unit 16 is attached to the flange portion. It is inserted and mounted until 35a hits. Further, by inserting and mounting the semiconductor radiation detector unit 16, a connector (not shown) of the camera 31 and the semiconductor radiation detector unit 16 is connected, and signals between the camera 31 and the semiconductor radiation detector unit 16 are connected. Line and power line are connected.

(Operation of PET system)
Next, the operation of this PET system will be described with reference to FIGS. 12 and 13 (refer to FIGS. 10 and 11 as appropriate).
The semiconductor radiation detector 13A outputs a detection signal indicating the energy level of incident radiation (γ rays, X-rays) according to the amount of electrons collected on the anode electrode A, that is, the magnitude of electric charges. Upon receiving this detection signal, the analog detection circuit 27 generates an analog peak value signal. The analog peak value signal is converted into a digital peak value signal by the ADC 25 and input to the digital ASIC 26A. On the other hand, a timing signal is output from the timing pick-off circuit 24a of the analog detection circuit 27 to the digital ASIC 26A (see FIG. 12).

In the digital ASIC 26A, the peak value of the detection signal, the detection time determined by the timing signal, and the ID of the semiconductor radiation detector uniquely identifying each semiconductor radiation detector 13A are used as a set signal. 10 and the subsequent data processing device 32 as shown in FIG. Therefore, in the data processing device 32 at the subsequent stage, first, based on the data transmitted from the digital ASIC 26A, first, the counting by the leakage characteristic X-ray is performed by the same method as described with reference to the flowchart of FIG. 6 in the first embodiment. A drop correction process is performed on a detection signal equal to or greater than a predetermined threshold (Eth), and it is determined whether or not γ-rays having incident radiation energy (E0) generated by electron pair annihilation have been detected.
At this time, when the semiconductor radiation detectors 13A are arranged in a three-dimensional manner, the second semiconductor radiation detectors 13A adjacent to the first semiconductor radiation detector 13A are adjacent to each other in a three-dimensional space. It is.

Next, in the simultaneous measurement device 32a, simultaneous measurement processing (processing that considers two γ-rays with a predetermined energy in a predetermined extremely short time window as two γ-rays with the same root) or scattered radiation processing ( When three or more gamma rays are detected in a predetermined very short time window, a process for correcting the influence of Compton scattering generated by collision between photons and electrons and gamma rays scattered in the body is performed.
In addition, the data processing device 32 performs accumulation of data, determination of a generation position of γ rays, generation of image data to be displayed on the monitor 33, and the like.
Note that the counting loss correction process and the simultaneous measurement process by the characteristic X-ray may be performed in parallel in the simultaneous measurement device 32a.

In the PET system 30, the data processing device 32 determines the generation position of γ-rays based on the accumulated data, and the functional image data, that is, fluorine 18 based on the determined generation position of γ-rays. Image data of the affected area of the subject in which ( ¹⁸ F) is concentrated is generated. In the PET system 30, the affected area of the subject is displayed on the monitor 33 (see FIG. 10) based on the functional image data output from the data processing device 32.

  Since the semiconductor radiation detection element 211 used in the second embodiment generates charges in response to light, the housing 35 is made of a light-shielding material such as aluminum or an aluminum alloy. It is configured so as to eliminate a gap through which light enters. That is, the housing 35 has a light shielding property. If the light shielding property is ensured by other means, the housing 35 itself does not need to have the light shielding property.

  As described above, according to the present embodiment, as in the first embodiment, correction of counting omission due to leakage characteristic X-rays is performed on a detection signal having a peak value exceeding the threshold Eth, and detection detection of normal incident radiation is performed. Therefore, a clear image can be obtained. Further, in FIG. 8, the imaging sensitivity can be improved only for the incident radiation energy peak 72 as compared with the case where the normal detection is determined by the energy window having a predetermined half width.

As is apparent from FIG. 8, the semiconductor radiation detector 13A mounted on the PET system 30 according to the second embodiment includes a semiconductor radiation detection element 211 manufactured from a single crystal of bromine-rich thallium bromide. Therefore, for example, energy peaks of ¹³⁷ Cs γ-rays appear remarkably at a position of 662 keV and a position of 589 keV, so that the energy resolution is excellent. Therefore, according to the PET system 30 according to the second embodiment, the detection accuracy of γ rays can be improved.

  Further, the semiconductor radiation detector 13A mounted on the PET system 30 according to the second embodiment has excellent withstand voltage characteristics. For example, in the semiconductor radiation detection element 211 having a thickness of 1.0 to 0.5 mm, in a conventional semiconductor radiation detection element using a single crystal of thallium bromide, the bias voltage that can be applied is about 400V. On the other hand, in the semiconductor radiation detection element 211 used in the second embodiment, a bias voltage of about 800 V can be applied. Therefore, according to the PET system 30 according to the second embodiment, since a high bias voltage can be applied to the semiconductor radiation detection element 211, charges can be collected efficiently, and the γ-ray detection signal The output level can be increased.

  Further, the semiconductor radiation detector 13A mounted on the PET system 30 according to the second embodiment has excellent noise characteristics, and so-called popcorn noise has occurred in the case of using a conventional semiconductor radiation detection element. On the other hand, popcorn noise does not occur in this semiconductor radiation detector 13A. Therefore, according to the PET system 30 according to the second embodiment, the S / N ratio of the γ-ray detection signal is further improved in combination with the increase in the output of the γ-ray detection signal. Further, since popcorn noise does not occur, a DC power source V applied between the cathode C and the anode A of the semiconductor radiation detector 13A when the popcorn noise exceeding the allowable level, which is necessary for the conventional PET system, is generated. A so-called reset circuit for blocking is not necessary in the PET system 30 according to the second embodiment.

  Further, in the PET system 30 according to the second embodiment, since the semiconductor radiation detection element 211 is composed of a single crystal of bromine-enriched thallium bromide, a conventional cadmium telluride single crystal is used. Compared with PET system, it is much cheaper.

  Furthermore, in the second embodiment, the PET system 30 (see FIG. 10) has been described as an example of a nuclear medicine diagnosis apparatus. However, the present invention is not limited to the PET system 30 but is applied to a SPECT (Single Photon Emission Computed Tomography) system. Can be applied. Incidentally, although the PET system and the SPECT system are common by taking a three-dimensional functional image of the human body, the SPECT system does not perform simultaneous measurement processing because the measurement principle is to detect single photons. For this reason, the collimator which regulates the incident position (angle) of a gamma ray is provided.

  The nuclear medicine diagnostic apparatus according to the present invention is inexpensive and improves energy resolution, withstand voltage characteristics, noise characteristics, and the like, and can be effectively used as a nuclear medicine diagnostic apparatus with clear images and good imaging performance. .

  In the first and second embodiments, the thallium bromide semiconductor radiation detection element has been described. However, the present invention is also applicable to other radiation detection elements that generate characteristic X-rays by incident radiation. it can.

1 is an overall configuration diagram of a γ camera imaging system according to a first embodiment of the present invention. It is an internal block diagram of the camera shown in FIG. FIG. 3 is a block diagram of a detector connector board and an ASIC module board shown in FIG. 2. It is a schematic diagram of a semiconductor radiation detector. (A), (b) is a schematic diagram which shows the behavior of the characteristic X-ray which an incident gamma ray produces | generates. It is a figure which shows the flowchart of a correction process. It is the figure which showed typically the positional relationship of the 1st semiconductor radiation detector which output the 1st detection signal, and the 2nd semiconductor radiation detector which outputs the 2nd adjacent detection signal. It is a spectrum figure of the peak value of the detection signal of the semiconductor radiation detector using the single crystal of bromine enrichment in the case where the radiation source is cesium 137 ( ¹³⁷ Cs). It is the figure which showed typically the imaging result before and behind performing a correction process, (a) performs the correction process of a detection radiation (gamma ray), (b) performs the correction process of a detection radiation (gamma ray). FIG. It is a perspective view which shows the structure of PET system as a nuclear medicine diagnostic apparatus concerning the 2nd Embodiment of this invention. It is the figure which showed typically the cross section of the camera in the PET system shown in FIG. It is a block diagram of the semiconductor radiation detector unit used for the PET system shown in FIG. It is a schematic diagram of the semiconductor radiation detector used for the semiconductor radiation detector unit shown in FIG. (A) is a partially broken perspective view of the camera showing a state when the semiconductor radiation detector unit is attached to the camera, and (b) is an AA cross-sectional view of the central portion of the camera.

Explanation of symbols

1 γ Camera Imaging System (Nuclear Medicine Diagnostic Equipment)
DESCRIPTION OF SYMBOLS 3 Camera 4 Camera stand 5a Data collection apparatus 5b Data analysis apparatus 7 Monitor 11 Collimator 11a Window part 12 Light shielding shield case 13, 13A Semiconductor radiation detector 15 Detector connector board 16 Semiconductor radiation detector unit 18 ASIC module board 22 Capacitor 23 Resistance 24 Analog Integrated Circuit 24a Timing Pickoff Circuit 24b Charge Amplifier 24c Polarity Amplifier 24d Band Pass Filter 24e Peak Hold Circuit (Peak Height Analysis Circuit)
25 Analog-to-digital converter 26 Digital integrated circuit (correction means)
26A Digital integrated circuit 27 Analog detection circuit (signal processing circuit)
30 PET system (nuclear medicine diagnostic equipment)
31 Camera 31a Lid 31b Unit support member 32 Data processing device (correction means)
32a Simultaneous measurement device (correction means)
33 Monitor 34 Bed 35 Case 35a Flange 211, 211a, 211b Semiconductor radiation detection element A Anode electrode C Cathode electrode C1 Connector C2 Output connector P Subject V DC power supply 41 γ-ray 61 Inside element 62, 65 Characteristic X-ray 64 Surface Neighborhood 80 First semiconductor radiation detector 81 Second semiconductor radiation detector

Claims (14)

A plurality of semiconductor radiation detectors are used to detect radiation emitted from the body of the subject or radiation that has passed through the subject's body, and a captured image of the subject is detected based on a detection signal of the radiation. A nuclear medicine diagnostic device,
Among the plurality of semiconductor radiation detectors, the detected radiation energy (E1) detected by the incident radiation of energy (E0) entering the first semiconductor radiation detector is the energy (E0) of the incident radiation, Determining whether the incident radiation is generated and equal to a difference from a characteristic X-ray energy (EX) detected in a second semiconductor radiation detector adjacent to the first semiconductor radiation detector; A nuclear medicine diagnosis apparatus comprising: a correction unit that determines whether to correct a detection signal of the first semiconductor radiation detector based on a determination result.   2. The nucleus according to claim 1, wherein the correction unit determines that incident radiation of the energy (E0) is incident on a position of the first semiconductor radiation detector when E1 = E0−EX. Medical diagnostic device. In a nuclear medicine diagnostic apparatus that detects radiation emitted from a radiation source inside or outside a subject and obtains a captured image of the subject based on a detection signal of the radiation.
A plurality of semiconductor radiation detectors comprising an element that generates characteristic X-rays upon incidence of the radiation;
A first determination as to whether adjacent semiconductor radiation detectors of the plurality of semiconductor radiation detectors have detected radiation substantially simultaneously; and
A second determination as to whether or not the energy of the radiation detected by any one of the adjacent semiconductor radiation detectors is equal to the energy (EX) of the characteristic X-ray;
Whether the sum of the energy of the radiation detected by both of the adjacent semiconductor radiation detectors is equal to the energy (E0) of the radiation emitted from the radiation source, or whether one of the adjacent semiconductor radiation detectors is A third determination is made as to whether or not the detected radiation energy is equal to the difference between the energy (E0) of the radiation emitted from the radiation source and the energy of the characteristic X-ray, and the radiation is determined based on these three determinations. A nuclear medicine diagnostic apparatus comprising a data processing device for determining whether or not detection is detected.   The nuclear medicine diagnosis according to claim 3, wherein it is determined that the radiation of the energy (E0) is incident on the semiconductor radiation detector having the higher radiation energy detected by the adjacent semiconductor radiation detector. apparatus.   2. The nuclear medicine diagnosis apparatus according to claim 1, wherein the plurality of semiconductor radiation detectors are arranged two-dimensionally or three-dimensionally.   6. The nuclear medicine diagnosis according to claim 5, wherein the semiconductor radiation detection element of the semiconductor radiation detector is configured using a single crystal of bromine-enriched material in which bromine is enriched in thallium bromide. apparatus.   The nuclear medicine diagnostic apparatus according to claim 6, wherein the bromine-enriched single crystal contains trivalent thallium. A DC power supply for applying a predetermined DC voltage between an anode electrode and a cathode electrode disposed on both sides of the semiconductor radiation detection element;
A signal processing circuit for processing the detection signal output from either the anode electrode or the cathode electrode;
The nuclear medicine diagnosis apparatus according to claim 6, further comprising:   The nuclear medicine diagnosis apparatus according to claim 8, wherein the signal processing circuit includes a pulse height analysis circuit that analyzes radiation energy.   The nuclear medicine diagnosis apparatus according to claim 1, wherein the semiconductor radiation detector is housed in a housing that shields light.   A radiation camera comprising: the semiconductor radiation detector used in the nuclear medicine diagnosis apparatus according to claim 1; and the correction unit.   A radiation camera comprising: the semiconductor radiation detector used in the nuclear medicine diagnosis apparatus according to claim 3; and the data processing apparatus. Nuclei that detect radiation emitted from the body of a subject or transmitted through the subject using a plurality of semiconductor radiation detectors and obtain a captured image of the subject generated based on a detection signal of the radiation A method for detecting radiation in a medical diagnostic device, comprising:
When it is known that the incident radiation incident on the semiconductor radiation detector has a predetermined energy (E0),
A first step in which the incident radiation is incident on a first semiconductor detector of the plurality of semiconductor detectors, and the first semiconductor detector detects a signal of detected radiation energy (E1);
Whether the detected radiation energy (E1) detected by the first semiconductor radiation detector is equal to the difference between the incident radiation energy (E0) and the characteristic X-ray energy (EX) generated by the incident radiation. A second step of determining;
Detected radiation energy (E2) detected by a second semiconductor radiation detector adjacent to the first semiconductor radiation detector causes the incident radiation to be generated in the first semiconductor radiation detector and the second semiconductor. A third step of checking whether the characteristic X-ray of the energy (EX) emitted to the radiation detector is detected;
The detected radiation energy (E1) detected by the first semiconductor radiation detector is E1 = E0−EX, and the detected radiation energy (E2) detected by the second semiconductor radiation detector substantially simultaneously is E2. A fourth step of determining that incident radiation of the energy (E0) has entered the position of the first semiconductor radiation detector when = EX;
The radiation detection method in the nuclear medicine diagnostic apparatus characterized by including this. Radiation detection in which a radiation detection determination device determines the presence or absence of radiation detection based on signals from a plurality of semiconductor radiation detectors containing elements that generate characteristic X-rays upon incidence of radiation emitted from a radiation source A method,
The radiation detection determination device includes:
Among the plurality of semiconductor radiation detectors, a determination as to whether adjacent semiconductor radiation detectors have detected radiation substantially simultaneously,
Determining whether the energy of the radiation detected by any one of the adjacent semiconductor radiation detectors is equal to the energy (EX) of the characteristic X-ray;
Whether the sum of the energy of the radiation detected by both of the adjacent semiconductor radiation detectors is equal to the energy (E0) of the radiation emitted from the radiation source, or whether one of the adjacent semiconductor radiation detectors is Determining whether the energy of the detected radiation is equal to the difference between the energy of the radiation emitted from the radiation source (E0) and the energy of the characteristic X-ray;
A radiation detection method, wherein the presence or absence of radiation detection is determined based on this determination.

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