JP2007114219A - Nuclear medicine imaging apparatus - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a nuclear medicine imaging apparatus having low noise in image quality and high spatial resolution. <P>SOLUTION: A γ-ray detection signal processor 41 includes a correction data generator 16 for outputting a time correction value Δτ and an absorption position N' of a γ-ray inside semiconductor detection elements 2a-2n, on the basis of a rising time of a signal output from semiconductor detectors 1a-1n by absorption of the γ-ray. A correction operation device 25 outputs correction time information (τ-Δτ) as a precise γ-ray absorption time, and high precision address information (N*N') in which the absorption position N' of the γ-ray inside the semiconductor detectors 1a-1n is added, respectively to a simultaneous counting device 26, with reference to the time correction value Δτ. <P>COPYRIGHT: (C)2007,JPO&INPIT

Description

  The present invention relates to nuclear medicine imaging in which a drug labeled with a radioisotope (RI) is administered to a subject such as a patient, and γ rays emitted from the RI are detected to obtain an RI distribution in the subject. Relates to the device.

  An apparatus for acquiring a RI distribution in a subject by introducing a radiopharmaceutical labeled with RI (hereinafter simply referred to as a drug) into a subject such as a patient and detecting γ-rays emitted from the RI. This is called a nuclear medicine imaging device (see, for example, Patent Document 1). Typical nuclear medicine imaging devices include a gamma camera, single photon emission CT (SPECT), positron emission CT (PET), and the like.

  A gamma camera is a device that measures gamma rays emitted from the inside of a subject with a flat detector and images the plane distribution. A collimator is attached to the front of the flat detector, and the incident direction of gamma rays is determined. Restrict and give directionality. The planar detector has a plurality of radiation detectors arranged in a plurality of columns and a plurality of rows.

  A single photon emission tomography apparatus (SPECT apparatus) detects a γ-ray emitted from the subject by arranging a planar detector similar to the above-described gamma camera around the subject, and X-ray CT. Is a device that obtains a tomographic image of a subject by imaging processing in the same manner as in FIG. In the SPECT apparatus, as in the gamma camera, a collimator is attached to the front surface of the flat detector to limit the incident direction of γ rays. The RI used in the SPECT apparatus is a nuclide that emits a single gamma ray (for example, 99mTc or 123I). The SPECT apparatus can image the distribution of at least one of the RIs in the subject and know organ circulation and metabolic information.

A proton emission tomography apparatus (PET apparatus) detects γ rays emitted from the inside of a subject by a ring detector arranged around the subject, and performs an imaging process based on the detection signal of the γ rays. This is an apparatus for obtaining a tomographic image showing the RI distribution in the subject. The ring detector has a number of radiation detectors arranged annularly and in the longitudinal direction of the bed. The PET apparatus administers a drug labeled with a nuclide that emits positrons (β ⁺ ) to a subject, and releases β ⁺ to combine with electrons and disappear, almost in the opposite direction (180 ± 0.6 ° ), A pair of γ-rays of 511 keV to be detected.

The PET apparatus can determine the incident direction of a pair of γ rays by selecting a pair of γ rays detected at the same timing by the coincidence counting device. For this reason, unlike the gamma camera and the SPECT apparatus, the PET apparatus does not need to use a collimator. Examples of positron emitting nuclides used in the PET apparatus include ¹⁸ F, ¹⁵ O, and ¹¹ C. For example, fluorodeoxyglucose (2- [F-18] fluoro-2-deoxy-D-glucose, ¹⁸ F-FDG), which is a drug labeled with 18F, is administered to a tumor tissue due to glucose metabolism when administered into a subject. It is used to identify the tumor site by utilizing the high accumulation in the tumor.

  By the way, in a conventional nuclear medicine imaging apparatus, a scintillator mainly composed of a substance such as bismuth germanium oxide (BGO) or thallium-added sodium iodide (NaI (Tl)) is used as a radiation detector for detecting γ rays. Used. In this radiation detector, incident γ-rays are once converted into weak light by a scintillator, and the weak light is converted into an electric signal by a photomultiplier tube or a photodiode. Therefore, there has been a problem that the nuclear medicine imaging apparatus is increased in size.

Therefore, attention is now focused on semiconductor radiation detectors (hereinafter referred to as semiconductor detectors) composed of semiconductor cells such as cadmium telluride (CdTe) and zinc cadmium telluride (CdZnTe) (see, for example, Patent Document 2). These semiconductor detectors directly convert gamma rays into charge carriers (electrons and holes). Therefore, since gamma rays can be detected by individual semiconductor cells, the apparatus can be expected to be smaller and lighter than when a scintillator and a photomultiplier tube are used. Also, the number of charge carriers generated is much larger than the number obtained by the scintillator detector, which means that good energy resolution can be obtained.
JP 11-337645 A (paragraph number [0002], FIG. 1) JP 2003-79614 A (paragraph number [0016], FIG. 1)

  An object of the present invention is to provide a nuclear medicine imaging apparatus capable of obtaining an image with less noise when a semiconductor radiation detector is used.

  In order to achieve the object, in the present invention, based on the rise time and the function of analyzing the rise time of the voltage signal (γ-ray detection signal) output from the semiconductor radiation detector corresponding to the induced charge amount, And a function of providing time correction information for accurately obtaining the absorption time in the semiconductor detection section of the detector.

  Preferably, the first absorption time information obtained corresponding to the reception timing of the γ-ray detection signal output from the radiation detector by the absorption of γ-rays by the semiconductor detection unit is obtained using the rise time of the γ-ray detection signal. The second absorption time information is generated by correcting with the obtained time correction information, and an image is generated with information obtained based on a plurality of pairs of γ-ray detection signals selected by the second absorption time information. . The information obtained on the basis of the plural pairs of γ-ray detection signals is obtained by counting two digital signals output and selected from the correction arithmetic unit as a specific example as one signal. Contains numeric values.

  Preferably, rise time-time correction value characteristic data (used in the time correction data generation device) for obtaining a correlation between the rise time of the voltage signal and the absorption time in the semiconductor detection element in advance, and the rise time of the voltage signal and the semiconductor Rise time-absorption position characteristic data (used in the position correction data generation apparatus) for obtaining the correlation with the absorption position in the detection element is created. A correction data generation device including a rise time analyzer, a time correction data generation device incorporating rise time-time correction value characteristic data, and a position correction data generation device incorporating rise time-absorption position characteristic data Incorporate into. This correction data generation apparatus is an apparatus in which a time correction data generation apparatus and a position specifying apparatus are integrated. A correction value based on the time correction data generation device is input to the time correction device, and a value obtained by subtracting the correction value from the detection time is output as an absorption time. Further, the detector address (absorption position in the detector) is output as new address data from the absorption position data based on the position correction data generation device.

  The above rise time-time correction value characteristic data and rise time-absorption position characteristic data are generated before the operation of the imaging apparatus, and thereafter, characteristic tests are periodically performed according to the time-dependent characteristics of the semiconductor radiation detector. The characteristic data can be updated. In addition, the number of the rise time sections and the absorption position sections in the characteristic data can be adjusted according to the desired resolution and image quality.

  One correction data generation apparatus may be provided for one semiconductor radiation detector, but one correction data generation apparatus may be provided for each of the plurality of semiconductor radiation detectors according to incident γ rays. In consideration of the characteristics of the individual semiconductor detection elements (semiconductor detection units), the characteristic data in the correction data generation apparatus may be set for each correction data generation apparatus, but there are several absorption position divisions. If it is, it is not necessary to set the characteristic data for each correction data generation device, and one typical characteristic data may be applied to a plurality or all of the correction data generation devices. Even in such a case, it is possible to sufficiently improve the time correction and the accuracy of the absorption position.

  According to the present invention, spatial resolution can be improved and an image with reduced noise can be obtained.

  The inventors discovered a new problem when examining a nuclear medicine imaging apparatus to which a semiconductor detector is applied in detail. A new problem in the semiconductor detector discovered by the inventors will be described in detail below. The phenomenon that occurs internally when γ rays are incident on the semiconductor detector will be described in detail with reference to FIGS.

  As shown in FIG. 10, the semiconductor detector 60 includes a semiconductor detection element 69 made of, for example, CdTe, and an anode 63 and a cathode 64 formed at both ends of the semiconductor detection element 69. Incident γ-ray 65 enters the semiconductor detector 60 and interacts with the semiconductor detection element 69 at different γ-ray absorption positions 66A and 66B inside the semiconductor detection element 69. Inside the semiconductor detection element 69, electrons 67 and holes 68 are generated by absorption of incident γ rays 65. Since a voltage is applied to the semiconductor detection element 69, the electrons 67 move toward the anode 63 and the holes 68 move toward the cathode 64.

  Due to the movement of these charge carriers (electrons 67, holes 68), charges are induced at both ends of the semiconductor detection element 69, and a voltage signal proportional to the induced charge amount is output by a subsequent amplifier (not shown). A signal processing device (not shown) in the subsequent stage recognizes the detection time and energy of the incident γ-ray 65 by detecting from this voltage signal.

FIGS. 11A and 11B show the time lapse of the output voltage signal at the γ-ray absorption positions 66A and 66B. Here, the time lapse of the voltage signal output from the amplifier is composed of two components, one is a high-speed component due to electrons, and the other is a low-speed component due to holes. These differences are caused by differences in mobility. For example, in CdTe, the mobility of electrons 67 is about 1100 cm ² / V / second, and the mobility of holes 68 is about 100 cm ² / V / second, which is about 1 Different digits.

  That is, this means that the time lapse of the voltage signal changes depending on the γ-ray absorption positions 66A and 66B in the semiconductor detection element 69. When γ-rays are absorbed at the γ-ray absorption position 66A in the vicinity of the cathode 64, charges are induced in a short time because the distance to the cathode 64 is short even with the low-speed hole component 32A. Also for the electronic component 31A, the distance to the anode 63 is long, but since the speed is high, charges are induced in a short time. As a result, the voltage signal 33A obtained by adding both the hole component 31A and the electron component 31B also rises at a high speed (see ΔT in FIG. 11A) and gradually attenuates after reaching the maximum value. .

  On the other hand, when γ rays are absorbed at the γ ray absorption position 66B in the vicinity of the anode 63, the electronic component 31B is induced in a very short time because the distance to the anode 63 is short. Since the hole component 32B has a long distance to the cathode 64 and a low speed, it rises slowly (see ΔT in FIG. 11B). Therefore, the time for the voltage signal 33B to reach the maximum value also becomes longer. Further, since the holes 68 are trapped at the time of movement to the cathode 64 and cannot be collected sufficiently, the amount of induced charges does not reach the amount of induced charge obtained at the absorption position 30A, causing a loss of signal amount. Thus, when a signal loss occurs depending on the γ-ray absorption position in the semiconductor detector, the image information of the patient created using the information obtained based on the γ-ray detection signal that is the output of the semiconductor detector. The inventors have found a new problem of noise.

As described above, the change of the output signal due to the difference between the γ-ray absorption positions 66A and 66B mainly depends on the hole which is a low speed component. For example, in CdTe, compared with Si (electron: 1500 cm ² / V / second, hole: 500 cm ² / V / second) and Ge (electron: 3900 cm ² / V / second, hole: 1900 cm ² / V / second). Since the mobility of both electrons and holes is slow and the difference in mobility between electrons and holes is large, the tendency to increase the time to reach the maximum value of the voltage signal 33B becomes prominent.

  The inventors have newly found that the above phenomenon occurring in the semiconductor detector causes the following loss for the nuclear medicine imaging apparatus. That is, for example, when detecting a pair of γ-rays emitted simultaneously in a PET apparatus and determining that the pair of γ-rays are incident at the same time, a time width called a coincidence time window is set in advance. It is determined that the two detection events obtained within are simultaneous. This time width is preferably as small as possible. Because if the time width is large, there is an opportunity not only to detect a pair of gamma rays within the time width (true coincidence), but also to detect gamma rays that have occurred in other parts of the body (incident coincidence) This is because it increases. The coincidence coincidence is a false phenomenon and causes noise on the image (tomographic image) of the subject.

  On the other hand, in the case of a scintillator detector, since the light emission phenomenon is used, the rising of the signal hardly changes no matter where the incident γ rays are absorbed in the detection element. The time difference between the line absorption time) and the time when the signal is detected by the subsequent timing detection device (detection time) hardly changes with respect to the absorbed position, and the signal rises quickly. Therefore, the scintillator detector is excellent in the simultaneous detection performance, the half width indicating the so-called coincidence resolution is short, about 10 [ns], and the time width can be set to approximately the same.

  However, in a semiconductor detector such as CdTe, the rise of the signal changes depending on the absorption position as described above, so that the variation in the γ-ray absorption time and the detection time becomes large. For this reason, it is necessary to increase the half-value width to about 20 [ns], and the time width called a coincidence time window must be set to be similarly large. Therefore, in order to improve the image quality of the image (tomographic image) of the subject when using a semiconductor detector, a coincidence resolution that is not inferior to that of a scintillator detector even when using a semiconductor detector is guaranteed. There is a need.

  The inventors have made the present invention in order to solve the newly discovered problem, that is, to reduce noise in an image. Specific embodiments of the present invention obtained through various studies in order to solve the new problem will be described below.

(First embodiment)
Next, a nuclear medicine imaging apparatus, specifically a PET apparatus, according to the first embodiment of the present invention will be described in detail with reference to FIGS. In the present embodiment, the same elements are denoted by lowercase letters after the numbers, and duplicate descriptions are omitted.

  First, characteristics indicating the relationship between the rise time ΔT and the detection time difference Δτ used in the time correction data generation device 17 and the position correction data generation device 18 (FIG. 5) of the PET apparatus of the present embodiment, and the rise time ΔT and absorption. A test apparatus A that obtains a characteristic indicating a relationship with the position Δx (N ′) will be described with reference to FIG. The test apparatus A includes semiconductor detectors 1a and 1b, a negative high-voltage power supply 5, preamplifiers 6a and 6b that amplify charges induced by the semiconductor detectors 1a and 1b, and outputs of the preamplifiers 6a and 6b. High-speed amplifiers 7a and 7b that shape and amplify the waveform at high speed, time pick-off devices 8a and 8b that detect timing of output signals from the high-speed amplifiers 7a and 7b, and output signals from the time pick-off device 8a are delayed by a predetermined time. The delay device 9, the time wave height converter 10 that outputs a pulse having a pulse height proportional to the time difference between the timing pulses output from the two time pick-off devices 8a and 8b, and the wave height that is proportional to the rise time of the output waveform of the high-speed amplifier 7a And a rise time analyzer 11 that outputs a pulse of The semiconductor detectors 1a and 1b include two semiconductor detection elements 2a and 2b of about 5 mm square, anodes 3a and 3b that are voltage application electrodes (anodes), and cathodes 4a and 4b that are cathodes. .

  Between the two semiconductor detectors 1a and 1b, a standard source 12 of a positron emitting nuclide such as 22Na and a lead collimator 13 having a hole diameter of about 100 μm are installed. First, the cathodes 4 a and 4 b of the semiconductor detectors 1 a and 1 b are arranged on the exit side of the collimator 13. Accordingly, the pair of γ rays 29 and 29 generated by the annihilation of the positrons emitted from the standard radiation source 12 enter the vicinity of the cathodes 4a and 4b of the semiconductor detectors 1a and 1b. Each charge induced by the absorption of γ rays in the semiconductor detection elements 2a and 2b is generated in the form of a waveform 36A as shown in FIG. The preamplifiers 6a and 6b to which the charges are input output a voltage signal having a waveform 37A. The high-speed amplifiers 7a and 7b to which this voltage signal is input output a voltage signal having a waveform 38A having a rising waveform 37A. The output signal of the high speed amplifier 7a is input to the time pick-off device 8a and the rise time analyzer 11. The output signal of the high speed amplifier 7b is input to the time pick-off device 8b. The time pick-off devices 8a and 8b output timing pulses by, for example, a constant fraction method. The time wave height converter 10 outputs a pulse having a wave height proportional to the detection time difference (time correction value) Δτ of both timing pulses.

  On the other hand, the rise time analyzer 11 outputs a pulse having a wave height proportional to the rise time ΔT that has elapsed from 10% to 90% of the input signal wave height, for example. In the current state (the state shown in FIG. 1), since the pair of γ rays 29 and 29 are absorbed in the vicinity of the cathodes 4a and 4b, as described above, the hole components generated inside the semiconductor detection elements 2a and 2b Voltage signals due to electronic components also rise at a high speed (see ΔT in FIG. 11A), and the time difference between the timing detections is small. That is, both the detection time difference Δτ and the rise time ΔT are small.

  Here, when one semiconductor detector 1a is moved and measured in the direction of arrow a in FIG. 1, the absorption position in the semiconductor detection element 2a is separated from the cathode 4a, so that the voltage signal generated in the semiconductor detector 1a. Rises slowly. Therefore, as shown in FIG. 4, the waveform of the induced charge output from the semiconductor detector 1a is 36B, the waveform of the voltage signal output from the preamplifier 6a is 37B, and the voltage signal output from the high speed amplifier 7a. The waveform is like 38B. Along with this, the detection time difference Δτ and the rise time ΔT increase. Thus, by moving the semiconductor detector 1a in the direction of the arrow a (FIG. 1) with respect to the collimator 13, the correlation between the rise time ΔT and the detection time difference Δτ as shown in FIG. A correlation with the movement amount Δx (hereinafter referred to as absorption position Δx) of the semiconductor detector 1a can be obtained.

  Accordingly, the rise time ΔT input based on the characteristics shown in FIG. 2 is converted into a detection time difference (time correction value) Δτ, and the obtained detection time difference Δτ is output, and FIG. Each of the position correction data generation devices that convert the rising time ΔT input based on the characteristics shown into the absorption position Δx and output the absorption position Δx (in the following embodiments, expressed as absorption position N ′ as position information), respectively. Can be created. As can be seen from FIGS. 2 and 3, the ΔT-Δτ characteristic and the ΔT-Δx (N ′) characteristic have distributions that vary, and it is difficult to convert them one-on-one. . Therefore, for example, the rising time ΔT is divided into three time sections with ΔT1 and ΔT2 as boundaries, and the time correction data generation apparatus corresponds to those time sections to Δτ1, Δτ2, Δτ3, and the position correction data generation apparatus May be converted into ΔN 1 ′, ΔN 2 ′, and ΔN 3 ′ in correspondence with the time intervals. Alternatively, a first relational expression between the rise time ΔT and the detection time difference Δτ and a second relational expression between the rise time ΔT and the absorption position N ′ are set for each of the three sections, and the time correction data generation apparatus The position correction data generation device may calculate the absorption position N ′ corresponding to the rise time ΔT using the second relational expression, using the first relational expression, and the detection time difference Δτ corresponding to the rise time ΔT.

  Next, FIG. 5 shows a PET apparatus (a kind of nuclear medicine imaging apparatus) according to the present embodiment using the ΔT-Δτ characteristic shown in FIG. 2 and the ΔT-Δx (N ′) characteristic shown in FIG. I will explain. First, a drug labeled with a positron emitting nuclide such as 18F is administered to the subject (examinee) H, and the system waits until the drug diffuses in the subject H and becomes ready for imaging. As a result, the drug accumulates in the affected area of the cancer in the subject. As shown in FIG. 5, the PET apparatus P has a configuration in which semiconductor detectors 1a, 1b... 1n made of, for example, CdTe are arranged in a ring shape. The bed B on which the subject H that is ready for imaging lies is inserted into the PET apparatus P. In this state, a large number of semiconductor detectors surround the subject H. These semiconductor detectors respectively detect a pair of γ rays emitted from the subject H in the opposite direction (180 ± 0.6 °) due to the drug. In this way, PET imaging is performed. The PET device P further includes a γ-ray detection signal processing device 41, a simultaneous coefficient device 26, and an image reconstruction device (image creation device) 28. A large number of semiconductor detectors installed in the PET apparatus P are divided into a plurality of groups in the circumferential direction of the PET apparatus P. The γ-ray detection signal processing device 41 is provided for each group.

  Hereinafter, one γ-ray detection signal processing device 41 will be described with reference to FIG. A voltage is applied from the high-voltage power supply 5 to the cathodes 4a, 4b,. When γ rays emitted from the subject H are incident on the semiconductor detection elements (semiconductor detection units) 2a, 2b,... 2n of each semiconductor detector 1, if the γ rays are absorbed by the semiconductor detection element 2, the absorption is performed. ... 3n and cathodes 4a, 4b... 4n are induced in amounts corresponding to the energy of the γ-rays generated. Each induced charge is output as a γ-ray detection signal to the preamplifiers 6a, 6b... 6n individually connected to the anodes 3a, 3b... 3n, and is amplified by the corresponding preamplifiers.

  Each of the low-speed amplifiers (proportional amplifiers) 14a, 14b,... 14n individually connected to the preamplifiers 6a, 6b... 6n corresponds to the preamplifiers 6a, 6b. The output from the preamplifier is input, and the voltage waveform 39A (or voltage waveform 39B) shown in FIG. 4 having a sufficient wave height corresponding to the above-described induced charge amount is output. The wave height of this voltage waveform reflects the γ-ray energy value. Each of the pulse height holding devices 15a, 15b,... 15n individually connected to the low-speed amplifiers 14a, 14b,... 14n receives the voltage waveform 39A (or voltage waveform 39B) from the corresponding low-speed amplifier, A pulse signal having a peak value of the voltage waveform is output to the pulse height discriminating devices 23a, 23b,... 23n of the signal processing device 20.

  On the other hand, the high-speed amplifiers 7a, 7b,... 7n individually connected to the preamplifiers 6a, 6b. The rising edge of the signal output from the amplifier is amplified at high speed, and a voltage signal having a waveform 38A (or waveform 38B) shown in FIG. 4 is output.

  The time pick-off devices (time detection devices) 8a, 8b,... 8n that input voltage signals from the corresponding high-speed amplifiers among the high-speed amplifiers 7a, 7b,. A timing pulse is output to the connected time discriminating device among the devices (time giving devices) 22a, 22b,... 22n according to the reception timing of the voltage signal. The time pick-off device and the time discriminating device constitute a time determining device. The voltage signals of the high-speed amplifiers 7a, 7b,... 7n are also input to the rise time analyzer (rise time generator) 11 of the correction data generator 16. The rise time analyzer 11 measures the rise time ΔT of each voltage signal based on each voltage signal from the high speed amplifiers 7a, 7b,... 7n, and the time correction data generation device 17 and the position correction data generation device. 18 is output. The correction data generation device 16 includes a rise time analyzer 11, a time correction data generation device 17 for obtaining a detection time difference Δτ based on the rise time ΔT output from the rise time analyzer 11, and an absorption position based on the rise time ΔT. A position correction data generation device 18 for obtaining N ′ is provided.

  The time correction data generation device 17 uses the time data characteristics (for example, a relational expression between the rising time ΔT and the detection time difference Δτ) obtained based on FIG. 2 to each of the high speed amplifiers 7a, 7b,. A detection time difference Δτ, which is a time correction value, is obtained for each rise time ΔT of the voltage signal, and is output to the correction arithmetic unit 25 of the signal processing device 20. These detection time differences Δτ are for γ-ray detection signals output from the semiconductor detectors 1a, 1b,. The position correction data generation device 18 uses the position data characteristic (for example, the relational expression between the rise time ΔT and the absorption position N ′) obtained based on FIG. 3 to detect the semiconductor detector for each of the rise times ΔT. The absorption positions N ′ of 1a, 1b... 1n are respectively obtained and output to the correction arithmetic unit 25 of the signal processing device 20.

  The signal processing device 20 performs the following processing. Each of the pulse height discriminating devices 23a, 23b,... 23n discriminates a pulse signal having a crest value equal to or greater than a threshold value and outputs it to the corresponding time discriminating device among the time discriminating devices 22a, 22b,. A clock signal that is a timing pulse from the high-speed clock device 19 is input to each of the time discriminating devices 22a, 22b,. Each time discriminating device converts a pulse signal having a peak value into a digital signal having time information of time τ corresponding to the clock signal when the timing pulse is input from the corresponding time pick-off device, and converts the digital signal to Output to the address discriminating device 24. The time discriminating devices 22a, 22b,... 22n and the address discriminating device 24 are included in the time / address judging device 21.

  The address discriminating device 24 is substantially connected to the time discriminating device that outputs the digital signal when the digital signal to which the time information of the time τ is given from each of the time discriminating devices 22a, 22b,. The address information N of the detected semiconductor detector is given to the digital signal. The address discriminating device 24 outputs the digital signal to which the time information τ and the address information N are added to the correction arithmetic device 25. The address information is given for each semiconductor detector provided in the PET apparatus P. In the correction arithmetic unit 25, time correction information (τ−Δτ) in which the time τ is corrected by the detection time difference (time correction value) Δτ is generated, and the absorption position N in the semiconductor detector 1 is further added to the address information N. Correction address information (N * N ') to which information' is added is generated. The correction arithmetic device 25 outputs a digital signal to which the time correction information (τ−Δτ) and the correction address information (N * N ′) are given to the coincidence counting device 26.

  In this embodiment, a preamplifier, a high-speed amplifier, a time pick-off device, a low-speed amplifier, a wave height holding device, a time discriminating device, and a wave height discriminating device are provided for each semiconductor detector. For this reason, the γ-ray detection signal, which is the output of one semiconductor detector, may change in signal form on the way, but is substantially input to these devices.

  In the present embodiment, the time correction device including the configuration for correcting the time τ in the time correction data generation device 17 and the correction calculation device 25 with the detected time difference Δτ, the position correction data generation device 18, and the correction calculation device. It can be said that a position correction device including a configuration for adding information on the absorption position N ′ in the semiconductor detector 1 to the address information N in 25 is provided. The corrected address information (N * N ′) includes the address information N (address of the semiconductor detector) of the corresponding semiconductor detector and the position N ′ where the gamma ray is absorbed in the semiconductor detector. The absorbed position is shown with high accuracy. Adding the information on the absorption position N ′ to the address information N corresponds to correction of the address information N.

  The configurations of the semiconductor detector 1 and the γ-ray detection signal processing device 41 shown in FIG. 6 are a part of those provided in the PET apparatus P. A large number of semiconductor detectors provided in the PET apparatus P are divided into a plurality of groups with the semiconductor detectors 1a, 1b... 1n as one unit. As shown in FIG. 5, these semiconductor detectors output γ-ray detection signals to γ-ray detection signal processing devices 41 that are different for each group. Each digital signal to which time correction information and correction address information are output, which is output from those γ-ray detection signal processing devices 41, is input to the coincidence counting device 26.

  The coincidence counter 26 performs coincidence using the digital signal output from the correction arithmetic unit 25 of each signal processor 20. That is, among the time correction information (τ−Δτ) given to these digital signals, two digital signals that fall within the time of the coincidence time window w [ns] are selected, and those digital signals are obtained by coincidence counting. Two digital signals are counted as one signal. The coincidence coefficient device 26 simultaneously counts each information of a combination of correction address information of two selected signals (a pair of correction address information) and a count value (number of occurrences of γ ray pairs) of the simultaneously counted signals. It memorize | stores in the memory (not shown) of the apparatus 26. FIG.

  The image reconstruction device 28 stores a plurality of pairs of correction address information and count values in the memory in the storage device 27. The pair of correction address information corresponds to the flight direction data of each of the pair of γ rays 29. The image reconstruction device 28 obtains the γ-ray generation density in each voxel by a known filter back projection method using a plurality of pairs of correction address information and count values. The image reconstruction device 28 creates a tomographic image at a radionuclide accumulation position, that is, a tumor position, based on the generation density of these γ-ray pairs. Information on this tomographic image is stored in the storage device 27. The tomographic image information is displayed on a display device (not shown).

  Note that the count value obtained by counting the two selected digital signals as one signal is information obtained based on a pair of γ-ray detection signals corresponding to the two selected digital signals.

  As described above, according to the present embodiment, the time correction value (detection time difference) Δτ obtained by the correction data generation device 16 can be obtained. The time close to the moment when is absorbed by the semiconductor detection element 2 can be determined as coincidence counting as time correction information (τ−Δτ). That is, since the time correction value Δτ is used, the coincidence time window w [ns] can be set short, and the coincidence coincidence is reduced. For this reason, it is possible to obtain a tomographic image with less noise than conventional PET apparatuses.

  In the present embodiment, for example, since the rise time ΔT is divided into three with ΔT1 and ΔT2 as a boundary, the coincidence resolution is the subsequent apparatus system (correction data generation device 16, signal processing device 20, coincidence counting device). 26, etc.) can be reduced to about half from the conventional 20 [ns], and performance comparable to about 10 [ns] of the scintillator can be exhibited. For this reason, the accuracy of the obtained tomographic image is further improved. Therefore, the width of the coincidence counting time window w can be set to the same level as that of the scintillator, and a highly accurate time resolution can be obtained.

  The address information is added to the address N of each semiconductor detector 1 and the gamma ray absorption position N ′ inside the address information is added, so that the position where the gamma ray is absorbed can be determined with higher accuracy. In other words, without changing the number of semiconductor detectors 1 and their associated devices, it is possible to effectively determine in which of the three regions the gamma rays are absorbed in the semiconductor detector in terms of measurement. This is equivalent to a case where the size of one semiconductor detector is reduced to one third and the number of semiconductor detectors is increased three times. In this embodiment, the spatial resolution on the image can be greatly improved as a result. Furthermore, the present embodiment in which one correction data generation device 16 is provided can simplify the configuration of the γ-ray detection signal processing device 41. The correction data generation device 16 may be installed for each radiation detector and connected to the corresponding high-speed amplifier.

  As described above, according to the present embodiment, the rise time of the voltage signal of γ-ray absorption absorbed by the subject is measured using the easily created rise time analyzer, and the time correction value and the absorption position information are added. By doing so, the coincidence resolution and the spatial resolution can be improved. As a result, significant noise reduction in image quality and improvement in spatial resolution can be realized.

(Second Embodiment)
A second embodiment of the present invention will be described with reference to FIG. A PET apparatus P1 which is a nuclear medicine imaging apparatus according to the present embodiment has a configuration in which the γ-ray detection signal processing apparatus 41 of the PET apparatus P is replaced with a γ-ray detection signal processing apparatus 51. In the present embodiment, the same components as those in the first embodiment are denoted by the same reference numerals, and the description thereof is omitted. The γ-ray detection signal processing device 51 includes a detection time correction device 35 instead of the correction data generation device 16. The detection time correction device 35 includes a time pick-off device 8, a rise time analyzer 11, a time correction data generation device 17, a delay device 52, and a pulse output timing adjustment device (hereinafter referred to as a timing adjustment device) 53, and a semiconductor detector. Provided for each. Each detection time correction device 35 is connected to a corresponding high-speed amplifier among the high-speed amplifiers 7a, 7b,. The position correction data generation device 18 inputs the rise time ΔT from the rise time analyzer 11 of the plurality of correction data generation devices 16A, calculates the absorption position N ′ in the semiconductor detection element 2, and signals the absorption position N ′. The data is output to the correction arithmetic unit 25 of the processing device 20.

  The voltage signals output from the high-speed amplifiers 7a, 7b,... 7n are respectively input to the time pick-off device 8 and the rise time analyzer 11 of the corresponding detection time correction device 35. The time pick-off device 8 outputs a timing pulse as described above. The timing pulse is delayed by a set time τd (hereinafter referred to as delay time τd) by the delay device 52 and then input to the timing adjustment device 53. The delay device 52 and the timing adjustment device 53 constitute a time adjustment device. The set time τd is set to the same value in all the delay devices 52. On the other hand, the rise time analyzer 11 measures the rise time ΔT of the voltage signal input from the corresponding high-speed amplifier and outputs it to the time correction data generation device 17 and the position correction data generation device 18.

  The time correction data generation device 17 outputs the obtained detection time difference Δτ to the timing adjustment device 53. The timing adjustment device 53 further delays the timing pulse delayed by the set time τd by (τs−Δτ) and outputs it to the time discriminating device 22. Here, τs is a value that ensures that τs> Δτ with respect to the assumed Δτ, and is set to a value common to all the timing adjustment devices 53. Further, as can be seen from the above, the delay time τd is a waiting time for processing from the measurement of the rising time ΔT by the rising analyzer 11 until the timing adjustment device 53 obtains (τs−Δτ). What is necessary is just to set to the time which can be performed. As a result, the time output from the timing adjustment device 53 of each correction data generation device 16A to the corresponding time determination device among the time determination devices 22a, 22b,... 22n is (τ + τd + τs−Δτ). This time is the time when the time correction information (τ−Δτ) in the first embodiment is delayed by a value (τd + τs) common to all the correction data generation devices 16A. Therefore, the time discriminators 22a, 22b,... 22n substantially acquire the time correction information (τ−Δτ) as in the first embodiment.

  On the other hand, the position correction data generation device 18 calculates the absorption position N ′ in the semiconductor detection element 2 from the position data characteristics obtained based on the characteristic diagram of FIG. As a result, as in the first embodiment, a digital signal to which time correction information (τ−Δτ) and correction address information (N * N ′) are added is output to the coincidence counting device 21.

  In the present embodiment, the effects produced in the first embodiment can be obtained except for the simplification of the configuration of the γ-ray detection signal processing device 41. Furthermore, in the present embodiment, the time at which the timing pulse received by the timing adjustment device 53 is output to the time determination device 22 is adjusted in advance in the correction data generation device 16. This eliminates the need for time correction and simplifies the logic design inside the signal processing device 20.

(Third embodiment)
Next, a third embodiment of the present invention will be described with reference to FIG. FIG. 8 is a block diagram showing a configuration of a gamma camera which is a kind of nuclear medicine imaging apparatus according to the third embodiment. In this embodiment, time information such as a time correction value is not particularly required, and thus no time correction data generation device is installed. However, a position correction data generation device is provided. In the present embodiment, the same components as those in the first embodiment are denoted by the same reference numerals, and the description thereof is omitted.

  A lead collimator 13 is installed closer to the subject H than the semiconductor detector 1. The collimator 13 forms a plurality of γ-ray passages 61 that are openings. The width of the semiconductor detector 1 is about 5 mm, and three γ-ray passages 61 are opposed to one semiconductor detector 1. From which opening 61 the γ-ray 29 enters the semiconductor detector 1, the rise time ΔT is measured by the rise time analyzer 11 that constitutes the subsequent position specifying device 62, and the position correction data generating device 18 is used as a basis. Further, it can be identified by obtaining the absorption position N ′ in the semiconductor detection element 2. For this reason, the position correction data generation device 18 has a rise time ΔT so that three absorption positions N ′ are determined at equal intervals corresponding to the arrangement of the three γ-ray paths facing one semiconductor detector 1. The interval between the sections is preset (see FIG. 3).

The operation of the gamma camera in this embodiment will be described.
First, a radiopharmaceutical containing a single photon emitting nuclide (for example, 99mTc) is administered to the subject H. This radiopharmaceutical is diffused in the subject H and waits until imaging becomes possible. The bed on which the subject H lies is moved to bring the subject H close to a gamma camera using the CdTe semiconductor detector 1 arranged in a plane. The high voltage power supply 5 is applied to the semiconductor detector 1 and the detection of the γ rays 29 is started. The γ rays 29 emitted from the subject H enter the semiconductor detection element 2 through the γ ray passage 61. When the γ-ray 29 is absorbed in the semiconductor detection element 2, an amount of charge (γ-ray detection signal) corresponding to the energy is induced. This charge is amplified by the preamplifier 6. The output signal from the preamplifier 6 is input to the low speed amplifier 14 and the high speed amplifier 7. The output signal amplified by the low-speed amplifier 14 is input to the wave height holding device 15, and a wave height pulse that is the output of the wave height holding device 15 is output to the wave height discriminating device 23 of the signal processing device 20. The output signal amplified by the high speed amplifier 7 is input to the rise time analyzer 11 of the position specifying device 26.

  The rise time analyzer 11 measures the rise time ΔT as in the first embodiment. The position correction data generation device 18 obtains absorption position information N ′ corresponding to the measured rise time ΔT, and outputs this absorption position information N ′ to the correction arithmetic device 25 of the signal processing device 20.

  The pulse height discriminating device 23 discriminates pulse height pulses having a peak value equal to or higher than a threshold value among the pulse height pulses output from the pulse height holding device 15. The address discriminating device 24 receives the pulse height pulse output from each wave height discriminating device 23 and discriminates the address N of the corresponding semiconductor detector 1. The address discriminating device 24 converts the input pulse height pulse into a digital signal to which the address N information is added, and outputs this digital signal. Then, the correction arithmetic unit 25 outputs a digital signal to which a correction address information (N * N ′) obtained by adding the information of the absorption position N ′ to the address N is added to a counting device (not shown). The counting device stores each count value obtained by counting each digital signal from the plurality of signal processing devices 20, that is, the correction arithmetic devices 25, and the corresponding correction address information in a memory (not shown) of the counting device. Remember. The image reconstruction device 28 generates a planar image for the subject H using the correction address information and the count value stored in the memory and the above-described filter back projection method. This image is displayed on a display device (not shown).

  As described above, according to the present embodiment, the collimator having the γ-ray passage 61 smaller than the width of the semiconductor detector 1 (the length in the patient's body axis direction) is obtained by obtaining the absorption position N ′ by the position specifying device 62. Even if 13 is disposed, it is possible to identify from which γ-ray passage 61 the γ-ray 29 is incident on one semiconductor detector 1, and as a result, the spatial resolution can be greatly improved. In addition, the gamma camera can be made compact without increasing the number of semiconductor detectors 1 and the number of devices associated therewith only by including the position specifying device 62 that can be easily created.

  This gamma camera can also be used as a SPECT device by rotating one or more planar semiconductor detectors around the subject.

(Fourth embodiment)
A fourth embodiment of the present invention will be described with reference to FIG. FIG. 9 is a block diagram showing a configuration of a gamma camera which is a kind of medical imaging apparatus according to the fourth embodiment. Also in the present embodiment, a time correction data generation device is not installed as in the third embodiment. In the present embodiment, the same components as those in the first embodiment are denoted by the same reference numerals, and the description thereof is omitted.

  Also in this embodiment, a lead collimator 13 is installed closer to the subject H than the semiconductor detector 1. One γ-ray passage 61 formed in the collimator 13 is opposed to one semiconductor detector 1 having a width of about 5 mm. The collimator 13 is inserted between the semiconductor detectors 1, and the width of the γ-ray passage 61 is equal to the width of the semiconductor detector 1. The anode 3 of the semiconductor detector 1 faces the γ-ray passage 61. The cathode 4 is provided in the semiconductor detection element 2 so as to be parallel to the anode 3. Due to the geometric shapes of the semiconductor detector 1 and the collimator 13, γ rays can be incident on the semiconductor detector 1 from a viewing angle represented by α.

  However, when the absorption region inside the semiconductor detection element 2 is divided into three absorption regions 71A, 71B, 71C and only the absorption region 71C farthest from the subject is selected, this effectively increases the height of the collimator 13. This means that the viewing angle of γ rays incident on the absorption region 71C is as small as β.

  In the case of imaging with the former large viewing angle α, the spatial resolution of the image obtained by the image reconstruction device 27 is rough, but the γ-ray count rate is high and inspection can be performed in a short time. This is called a normal inspection mode (short time simple mode). On the other hand, since only the event absorbed in the absorption region 71C is selected at the latter small viewing angle β, an image with a low count rate and a long inspection time but excellent spatial resolution can be obtained. This is called a high resolution inspection mode. In the present embodiment, the examination mode can be changed according to the demand of the examiner who is the subject H. For example, even if the resolution is low, it is used in the normal inspection mode when a short inspection is required, and in the high resolution inspection mode when a higher resolution inspection is required.

  The setting of the inspection mode, that is, the normal inspection mode and the high resolution inspection mode can be selected in advance by the inspector on the input screen of the controller 72 or the like. In response to the operation setting signal, the controller 72 determines the address of the corresponding semiconductor detector when the gamma ray is incident on one semiconductor detector regardless of which of the three absorption regions is absorbed in the normal inspection mode. In order to output a digital signal with N information added, in the high-resolution inspection mode, an address discrimination is performed so that a digital signal with address N information is output only when a permission signal from the position specifying device 73 is input. The device 24 is set up. When the normal inspection mode is selected, it is only necessary to identify the address N of the semiconductor detector 1, so that it is not necessary to activate the position specifying device 73. In the normal inspection mode, the controller 72 transmits a stop command to the position specifying device 73 and prevents the output of the permission signal from the position specifying device 73. In the normal inspection mode, as in the third embodiment, the counting device (not shown) counts the digital signal output from the address discriminating device 24 with the address N information added, and the counted value and address. N is stored in a memory (not shown) of the counting device. The image reconstruction device 28 performs reconstruction processing in the same manner as in the third embodiment using the address N and the count value input to the memory, and generates a planar image of the radionuclide accumulation position. This image is displayed on a display device (not shown).

  On the other hand, the controller 72 transmits a start command when starting the high-resolution inspection mode and a stop command when ending the high-definition inspection mode to the position specifying device 73, respectively. The position specifying device 73 outputs a permission signal to the address discriminating device 24 when a gamma ray is absorbed in the absorption region 71C while inputting a start command and a stop command. That is, the position specifying device 73 transmits a permission signal to the address discriminating device 24 when the rising time ΔT calculated by the rising time analyzer 11 is a value corresponding to the absorption region 71C. Even when the start command is input and the stop command is input, if the gamma rays are absorbed by the absorption regions 71A and 71B, the position specifying device 73 does not output a permission signal. Only when receiving the permission signal, the address discriminating device 24 outputs a digital signal to which the information of the address N is added to the counting device. Even in the high-resolution examination mode, the image reconstruction device 27 generates the planar image information of the subject H as described above.

  Further, the controller 72 always activates the position specifying device 73 as another mode so as to output the absorption position N ′ to the address discriminating device 24. Then, the address discrimination device 24 outputs the corrected address information (N * N ′) to the image reconstruction device 27. The image reconstructing device 28 can obtain images having different resolutions by selecting the correction address information (N * N ′).

  According to the present embodiment, the inspector can select the desired inspection time and spatial resolution and perform the inspection without replacing the heavy collimator 13. In the present embodiment, similarly to the third embodiment, one or more planar semiconductor detectors can be used as a SPECT apparatus by rotating around a subject.

  In each of the above embodiments, the detection time difference (time correction value) and the absorption position have been described as being divided into three, but the number is not limited to three. If the number of divisions is plural, it is sufficiently effective, and if the number is increased, the time resolution and the spatial resolution can be further improved. In each of the above-described embodiments, the semiconductor detector is described as being configured using cadmium telluride (CdTe). However, for example, cadmium zinc telluride (CdZnTe), gallium arsenide (GaAs), instead of CdTe, You may comprise a semiconductor detector using compound semiconductors, such as thallium bromide (TlBr).

  In the first embodiment, the time correction value has been described as being obtained from the characteristic diagram of FIG. 2, but the present invention is not limited to this, and the time correction value is determined by the γ-ray source to be imaged. It can be arbitrarily set according to the imaging parameters such as the type, the part to be imaged, and the imaging time.

It is a block diagram which shows the test apparatus for producing the time data and position data which concern on the 1st Embodiment of this invention. It is a characteristic view which shows the time data which concern on 1st Embodiment. It is a characteristic view which shows the position data which concerns on 1st Embodiment. It is a characteristic view which shows the output waveform in the main apparatus elements which concern on 1st Embodiment. 1 is an overall configuration diagram showing a PET apparatus which is a kind of nuclear medicine imaging apparatus according to a first embodiment. It is a block diagram which shows the gamma ray detection signal process part of the PET apparatus which concerns on 1st Embodiment. It is a block diagram which shows the gamma ray detection signal process part of the PET apparatus which concerns on the 2nd Embodiment of this invention. It is a block diagram which shows the structure of the gamma camera which is 1 type of the nuclear medicine imaging device which concerns on the 3rd Embodiment of this invention. It is a block diagram which shows the structure of the gamma camera which is 1 type of the nuclear medicine imaging device which concerns on the 4th Embodiment of this invention. It is explanatory drawing which shows the phenomenon which generate | occur | produces internally when a gamma ray injects into a semiconductor detector. FIG. 6 is a characteristic diagram showing the time lapse of the output voltage signal at the γ-ray absorption position, where (a) shows the time lapse of the output voltage signal at the γ-ray absorption position near the cathode, and (b) is the γ-ray absorption position near the anode. It is a figure which shows the time passage of the output voltage signal in.

Explanation of symbols

1, 1a, 1b, 1n Semiconductor detector 2, 2a, 2b, 2n Semiconductor detection element 3, 3a, 3b, 3n Anode 4, 4a, 4b, 4n Cathode 6, 6a, 6b, 6n Preamplifiers 7, 7a, 7b, 7n High-speed amplifiers 8, 8a, 8b, 8n Time pick-off device 9 Delay device 10 Time wave height converter 11 Rise time analyzer 12 Standard source 13 Collimator 14, 14a, 14b, 14n Low-speed amplifier 15, 15a, 15b, 15c Wave height holding device 16 Correction data generation device 17 Time correction data generation device 18 Position correction data generation device 19 High-speed clock device 20 Signal processing device 26 Simultaneous counting device 27 Storage device 28 Image reconstruction device 62, 73 Position identification device 29 γ ray P PET equipment

Claims (1)

  In a nuclear medicine imaging apparatus including a plurality of radiation detectors including a semiconductor detection unit on which γ rays are incident, receiving a γ-ray detection signal output from the radiation detector by absorption of the γ rays by the semiconductor detection unit The second absorption time information is generated by correcting the first absorption time information obtained corresponding to the timing with the time correction information obtained using the rising time of the γ-ray detection signal, and the second absorption time information is generated. A nuclear medicine imaging apparatus, wherein an image is generated based on information obtained based on a plurality of pairs of γ-ray detection signals selected by information.

Images