JP2015152356A - Dark countless radiation detection energy discrimination imaging system - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide a dark countless radiation detection energy discrimination imaging system which uses a combination of LSO and MPPC to convert X rays for energy discrimination imaging.SOLUTION: The present invention includes an X ray tube 6, LSO-MPPC radiation detection means 1 which contains a lamination body of LSO and MPPC formed on a ceramic substrate, being sealed up in an Al cap, for receiving X rays that penetrate a subject, a high speed current/voltage amplifier 2 which amplifies an optical current generated by the detection means to generate an event pulse, a high-speed comparator 3 for setting a lower limit wave height value of the event pulse, a counter card 9 which detects an event pulse of a lower limit voltage or higher for counting a square wave pulse, and a PC5 which re-configures a tomogram from the projection data obtained by counting the square wave pulses.

Description

  The present invention relates to a dark countless radiation detection energy discrimination imaging system using radiation detection means in which a scintillator having a short light emission lifetime and a multi-pixel photon counter (MPPC) are combined.

  In the classical model of an atom, electrons are rotating around a shell or orbit around a nucleus of positively charged protons and uncharged neutrons. The innermost shell follows the K shell, and then outwards, the L shell and the M shell. When the atoms are irradiated, electrons are emitted from the atoms by particles such as X-ray photons and electrons having sufficient energy. There are vacancies in the K shell, making the atoms unstable and high energy. Since the atoms try to restore the original configuration, for example, electrons in the outer shell such as the L shell are transitioned to the K shell. Since the electrons in the L shell are higher in energy than the electrons in the K shell, excess energy is emitted as X-rays when the electrons in the L shell transition to the K shell.

  Radiation detectors are scintillators that receive radiation such as X-rays and γ-rays and convert them into visible light, photomultiplier tubes (PMT) and photodiodes that detect visible light that has been converted and transmitted by the scintillator and convert it into electrical signals. (PD) and the like, a reflecting member for efficiently guiding scintillation light emitted from the scintillator crystal to the photodetector, and the like, and is composed of positron emission tomography (PET), scintillation camera ( SPECT: Single Photon Emission CT), energy discrimination X-ray CT system, spectrometer or dosimeter.

  PET, which plays an important role in the field of nuclear medicine, such as research on biological functions and higher-order brain functions, detects a pair of gamma rays using a simultaneous counting method using multiple radiation detectors, and accumulates simultaneous counting information. A histogram is created, and an image representing the spatial distribution of the frequency of occurrence of a pair of gamma rays in the measurement space is reconstructed based on the histogram.

  So far, energy discrimination X-ray image processing has been performed using monochromatic X-rays, and a very clean monochromatic parallel beam is generated using a synchrotron and a silicon (Si) single crystal. These parallel beams with a photon energy of about 35 keV can effectively absorb iodine density because X-ray photons with energy exceeding the K-edge energy of iodine (I) of 33.2 keV are effectively absorbed by iodine atoms. It is applied to fluorescent X-ray CT (XRF-CT) to be measured. It is also used for iodine (I) K-edge angiography, and the coronary artery is clearly imaged. However, it has been difficult to obtain sufficient machine time for monochromatic X-ray image processing. Therefore, the present inventors have obtained a cerium (Ce) X-ray generator (Ce) to realize IK edge angiography and XRF-CT since Ce K-series characteristic X-rays are effectively absorbed by iodine atoms. X-ray tube unit) was developed.

  In recent years, in response to the demand for higher performance of PET, development of scintillator crystals using new materials, development of new photomultiplier tubes and semiconductor photodetectors, parallel use of computers, development of new image reconstruction algorithms Etc. are becoming active. Recently, a high-speed photomultiplier tube (photomultiplier) PMT (Photo-Multiplier Tube) for TOF-PET (Time Of Flight-PET) using a time-of-flight difference of radiation, a flat panel position detection type PMT (PS-PMT) Avalanche photodiode (APD) arrays, Geiger mode multi-pixel APD (MPPC: Multipixel Photon Counter), etc. have been researched and developed, and the realization of various high-performance PET taking advantage of these devices is desired. ing.

  Generally, in MPPC and APD used in the Geiger mode, thermally generated dark current carriers are multiplied to generate pulses. This pulse is a dark pulse and causes a detection error. The number of dark pulses per second is the dark count rate [unit: cps (counts per second)]. Increasing the reverse voltage increases the detection efficiency but also increases the dark count accordingly. The dark count decreases as the temperature decreases. MPPC is a new type of photon counting device that consists of a plurality of Geiger mode APD pixels and can be used at room temperature. It is a small optical semiconductor device that has excellent photon counting capability and is not affected by a magnetic field at low voltage operation.

  Since MPPC is generally used in Geiger mode (see, for example, Patent Documents 1 and 2), a dark count of about 1 Mcps occurs, making it difficult to measure the spectrum of radiation (X-rays and γ-rays). In particular, since the peak value of the dark count pulse in MPPC is almost the same as that of low energy X-rays, it is necessary to reduce the dark count rate to about 0 cps in order to measure the spectrum of radiation.

  In PET, a lutetium oxyorthosilicate photomultiplier tube (LSO-PMT) is used (see, for example, Patent Document 3). Although dark count occurs, the pulse width is extremely short. Therefore, it is possible to adopt a method in which the dark count is not detected by increasing the time constant of the amplifier.

JP 2013-16637 A JP 2013-16638 A Special table 2012-503190 gazette

  However, since MPPC has a longer width than PMT, it is necessary to set the dark count rate to 0 cps by lowering the bias voltage.

  With recent advances in detector technology, cadmium telluride (CdTe) detectors have been widely developed and applied to gamma and X-ray spectral measurements. With these detectors it is possible to perform monochromatic image processing by sorting out appropriate photon energy ranges, both so-called levels and widths. And pixelated CdTe and cadmium zinc telluride (CZT) detector arrays have been developed and applied to preclinical energy discriminating (ED) X-ray CT systems.

  On the other hand, the inventors have used several ED (energy dispersive) -CT systems using CdTe detectors to achieve K-edge CT using iodine (I) and gadolinium (Gd) media. Developed. Under ideal conditions, the CdTe detector count rate has been reported to increase to 6 Mcps. However, when a general-purpose CdTe detector is used, it is difficult to prevent event pulse pileup in a region where the maximum count rate is 0.1 Mcps or more.

  The present inventors have developed a high-speed X-ray photon counting CT (PC-CT) system with a low energy resolution using an MPPC module and a short delay time scintillator in order to increase the count rate exceeding 1 Mcps.

Then, using a 1 × 1 × zinc oxide 1 mm ³ (ZnO) single crystal, the count rate is increased to 15Mcps, energy discrimination by choosing a lower level voltage of the event pulses shown from MPPC module (ED) The effect was confirmed. However, the dark count rate of the MPPC module is 1 Mcps, and it is difficult to measure the X-ray spectrum.

  Therefore, the present invention has been made in view of the above-mentioned problems of the prior art, and an object of the present invention is to detect X-ray and γ-ray by combining a scintillator having a short light emission lifetime and a multi-pixel photon counter (MPPC). By applying a bias voltage immediately before Geiger mode to MPPC, it realizes a high count rate without dark count without greatly reducing the sensitivity of MPPC. An object is to provide a dark countless radiation detector-applied energy discrimination imaging system that is excellent in conversion efficiency and can reduce the amount of X-ray exposure to a subject.

  In order to achieve the above object, a dark countless radiation detection energy discrimination imaging system according to an embodiment of the present invention is selectively set according to the characteristics of an object by a controller and emits a low dose rate X-ray. A multi-pixel photon counter (hereinafter referred to as MPPC) that is enclosed in an aluminum (Al) cap having a tube unit and a thin film window that serves as an X-ray photocathode, is formed on a ceramic substrate and is connected to a pair of electrodes. LSO-MPPC radiation (X-rays, γ-rays) comprising a Lu2 (SiO4) O (hereinafter referred to as LSO) single crystal laminated thereon and receiving X-rays emitted from the X-ray tube unit and passing through the subject. ) Amplifying the photocurrent generated from the detection means and the LSO-MPPC radiation detection means, A high-speed current / voltage amplifier for generating a pulse, a multi-channel analyzer (hereinafter referred to as MCA) for analyzing a pulse height of an event pulse from the high-speed current / voltage amplifier, and a photon number of the event pulse subjected to the pulse height analysis from the MCA And an X-ray photon counting and spectrum measurement system comprising a computer unit (hereinafter referred to as PC) expressed as a change due to a channel (hereinafter referred to as ch).

  A dark countless radiation detection energy discrimination imaging system according to another embodiment of the present invention is an X-ray tube unit that emits low-dose-rate X-rays selectively set according to the characteristics of a subject by a controller; The X-ray tube unit comprising an LSO single crystal enclosed in an Al cap having a thin film window serving as an X-ray photocathode and formed on a ceramic substrate and laminated on an MPPC to which a pair of electrodes are connected LSO-MPPC radiation detection means for receiving X-rays radiated from the subject and passing through the subject; a high-speed current / voltage amplifier for amplifying a photocurrent generated from the LSO-MPPC radiation detection means and generating an event pulse; A high-speed comparator for setting a lower limit peak value of an event pulse from the high-speed current / voltage amplifier; A tomographic image is reconstructed from a counter card that detects an event pulse exceeding the lower limit voltage generated from the high-speed comparator and counts square wave pulses, and projection data obtained by counting the square wave pulses by the counter card. And a PC.

  A dark countless radiation detection energy discrimination imaging system according to still another embodiment of the present invention includes an X-ray tube unit that emits low-dose-rate X-rays that is selectively set according to the characteristics of an object by a controller, An LSO single crystal encapsulated in an Al cap having a thin film window that becomes a photocathode of a line, formed on a ceramic substrate and stacked on an MPPC to which a pair of electrodes are connected, is provided from the X-ray tube unit. LSO-MPPC radiation detection means for receiving X-rays emitted and transmitted through the subject; a high-speed current / voltage amplifier for amplifying the photocurrent generated from the LSO-MPPC radiation detection means and generating an event pulse; The subject mounted on the turntable that rotates at the set rotation step and the high-speed current A scan stage that linearly scans the subject is set in the LSO-MPPC radiation detection means that is arranged between the pressure amplifier and attached to the center, and a lower limit peak value of the event pulse from the high-speed current / voltage amplifier is set. A high-speed comparator, a counter card for detecting event pulses exceeding the lower limit voltage generated from the high-speed comparator, counting square wave pulses, and linear scanning and rotation of the subject are repeatedly controlled by a two-axis controller. And a PC for reconstructing a tomographic image from projection data obtained via a counter card.

  Further, the LSO-MPPC radiation detection means includes an Al cap having a thin film window that shields and shields X-rays, and the LSO unit sealed in the Al cap and attached to the photoelectric surface. A laminated body of crystal and MPPC, and a BNC connector that accommodates the Al cap therein, receives X-rays that pass through the subject through the thin film window, and receives X-rays from the front surface of the LSO single crystal X-ray photons that detect (count) photons and then pass through MPPC are Compton scattered by the ceramic substrate behind to generate scattered rays and are converted into fluorescent rays.

  However, APD in MPPC can hardly detect X-rays.

  The X-ray tube unit generates X-rays at a low dose rate by reducing the X-tube current in a range of 1.0 μA to 2.0 mA order.

  Further, the high-speed current / voltage amplifier is connected to the MPPC of the LSO-MPPC radiation detection means, and has a first resistor loaded with the first bias voltage immediately before the Geiger mode, and a second bias lower than the first bias voltage. A high-speed operational amplifier (single power supply voltage feedback amplifier) loaded with a voltage, a micro capacitor and a second resistor connected between the MPPC and the high-speed operational amplifier, and a third resistor that bypasses the high-speed operational amplifier. And an inverting current / voltage amplifier circuit.

  The inventors of the present invention have a very low amount of light emitted from a Zno crystal having a decay time of 1 ns, so that LSO-MPPC radiation (MPSO and Lu2 (SiO4) O (hereinafter referred to as LSO) crystal having a decay time of 40 ns is used. The X-ray, γ-ray) detecting means is useful for high-speed X-ray photon counting and energy dispersion systems.

  In applying the LSO-MPPC detection means to the detector array of the PC-CT system, it is necessary to measure a spectrum using a multichannel analyzer (MCA) for event pulse wave height analysis. Since the maximum count rate of the conventional MCA is about 10 kcps, it is necessary to minimize the MPPC dark count rate.

  The LSO crystal has a specific gravity of 7.35 and easily counts 0.51 MeV γ-ray photons generated by annihilation. Since the LSO-MPPC detector will be applied to PET in the near future, it is necessary to measure the detector characteristics immediately before the Geiger mode in order to realize a high-speed noise reduction ED-CT system in medical images.

  According to the present invention, the LSO-MPPC radiation detection means for detecting radiation such as X-rays and γ-rays, which is a combination of an LSO single crystal, which is a scintillator with a short emission lifetime, and MPPC is used, and a bias voltage immediately before Geiger mode is applied to the MPPC. By applying, a high count rate without dark count can be realized without greatly reducing the sensitivity of MPPC. Therefore, using LSO-MPPC radiation detection means, X-ray photons are counted or converted, image processing is performed, photon energy or average energy is selected, and energy discrimination imaging is performed. There is an effect that it is possible to provide a dark countless radiation detection energy discriminating imaging system that is excellent in conversion efficiency into an electric signal and can reduce the X-ray exposure dose to the subject.

It is a block conceptual diagram for explaining conceptually the X-ray photon counting using the LSO-MPPC radiation detection means 1 of one embodiment of the present invention and the spectrum measurement system 30 using MCA. (A) is a conceptual vertical cross-sectional view for conceptually explaining the LSO-MPPC radiation detection means 1 of FIG. 1, (b) is a photograph of the appearance of the BNC connector 15, (c) is aluminum (Al) not shown. It is an external appearance photograph in the state where BNC connector 15 was stored in Al case main part 16b where a cap was removed. FIG. 2 is a conceptual block diagram for conceptually explaining a high-speed current / voltage amplifier 2 using the high-speed operational amplifier of FIG. 1. (A) is the measurement graph of the dark count rate measured using the LSO-MPPC radiation detection means 1 of FIG. 1, (b) is the graph which expanded the A section of (a). It is an X-ray spectrum graph measured using a CdTe detector. (A) is a gamma ray spectrum graph generated from Americium-241 ( ²⁴¹ Am) measured using a CdTe detector, and (b) is a gamma ray event pulse wave height analysis graph. It is a X-ray tube voltage change graph of the X-ray spectrum measured using the LSO-MPPC radiation detection means 1 of FIG. 1 shows a concept of a dark countless radiation detection energy discrimination imaging system (hereinafter, abbreviated as a high-speed X-ray photon counting CT system or a PC-CT system) 40 using the LSO-MPPC radiation detection means 1 of one embodiment of the present invention. It is a block diagram. CT imaging images by the PC-CT system 40 of FIG. 8 of iodine contrast agents of different concentrations [(a) 30 mg / ml I, (b) 15 mg / ml I] contained in two 10.5 mm diameter polyethylene vials. (Tomographic image), (A) is the lower limit voltage Vl of event pulse Vl = 0.5V, X-ray tube voltage V = 80kV, (B) is Vl = 1.5V, V = 80kV, (C) is Vl = 2. .5V, V = 80 kV, (D) is an external view of the polyethylene vial. CT imaging image (tomographic tomography) of the iodine contrast agent having different concentrations ((a) 15 mg / ml I, (b) 30 mg / ml I) in the acrylic phantom with two holes by the PC-CT system 40 in FIG. (A) is the lower limit voltage of event pulse Vl = 0.5V, X-ray tube voltage V = 80kV, (B) is Vl = 1.5V, V = 80kV, (C) is Vl = 2.5V , V = 80 kV, (D) is an external view of the acrylic phantom. FIG. 8A is a CT image (tomographic image) of the heart of a dog with iodine microparticles (iodine contrast medium) having a diameter of 15 μm in the coronary artery. FIG. 8A shows the lower limit voltage Vl of the event pulse. = 0.5V, X-ray tube voltage V = 80kV, (B) is Vl = 1.5V, V = 80kV, (C) is Vl = 2.5V, V = 80kV, (D) is the key to (D) FIG.

Hereinafter, a specific example of an embodiment for implementing an energy discrimination imaging system applied with a dark countless radiation detector of the present invention will be described with reference to the accompanying drawings.
Example 1

  As shown in FIG. 1, the dark countless radiation detector applied energy discrimination imaging system according to an embodiment of the present invention is selectively set by a controller 7 in accordance with the characteristics of a subject (not shown). An X-ray tube unit (hereinafter referred to as X-ray tube) 6 that emits photons (hereinafter referred to as X-rays) Xr, a multi-pixel photon counter (MPPC) 11 and a Lu2 (SiO4) O (hereinafter referred to as LSO) crystal 10. LSO-MPPC radiation (X-ray, γ-ray) detection means 1 that includes a laminate (see FIG. 2) and receives X-rays Xr emitted from the X-ray tube unit 6 and transmitted through the subject, and LSO-MPPC radiations A high-speed current / voltage amplifier 2 that amplifies the pulsed photocurrent generated from the detection means 1 and generates an event pulse, and a high-speed current / voltage amplifier A multi-channel analyzer (MCA) 4 for analyzing the pulse height of the event pulse from the computer, and a computer unit (PC) 5 for representing the event pulse subjected to the pulse height analysis from the MCA 4 as a change in the channel (ch) of the number of photons. A line photon counting and spectral measurement system 30 is included.

As shown in FIG. 2, the LSO-MPPC radiation detection means 1 of Example 1 of the present invention includes an aluminum (Al) cap 16 having a thin film window 16a that shields light and shields and serves as an X-ray photocathode. A laminated body of the LSO single crystal 10 and the MPPC 11 enclosed in the cap 16 and attached to the photocathode 16a, and a BNC connector 15 that accommodates the Al cap 16 therein are provided. A 1 × 1 × 1 mm ³ LSO single crystal 10 formed on the ceramic substrate 12 and laminated on the MPPC 11 to which a pair of electrodes 14 are connected is attached to a 1 × 1 mm ² photocathode and has a thickness of 0. It is sealed in an Al cap 16 having a 2 mm thin film window 16a to be shielded from light. Since the pulse peak photocurrent of the LSO-MPPC radiation detection means 1 is on the order of μA, an electromagnetic shield is not particularly required. Shielding is performed using the case 16b. The upper and lower surfaces of the ceramic substrate 12 in the Al cap 16 are insulated by black rubbers 13 and 13.

  The X-ray photon Xr transmitted through the subject is received through the thin film window 16a, the X-ray photon Xr is detected (counted) from the front surface of the LSO single crystal 10, and then the X-ray photon Xr transmitted through the MPPC 11 is the ceramic substrate behind 12 is Compton scattered and generates scattered radiation and converted to fluorescent radiation

  The X-ray tube unit 6 has been developed to generate an X-ray beam having a low dose rate by reducing the X-ray tube current in the range of 1.0 μA to 2.0 mA. X-ray photons Xr generated from the X-ray tube unit 6 are converted into scintillation photons by the LSO single crystal 10 of the LSO-MPPC radiation detection means 1 1.0 m away from the X-ray tube unit 6 and detected by the MPPC 11. The pulsed photocurrent flowing through the MPPC 11 is amplified by the high-speed current / voltage amplifier 2 to generate an event pulse. The event pulse from the high-speed current / voltage amplifier 2 is sent to an MCA (MCA4000 of PGT) 4 and changes in the X-ray tube voltage V and aluminum (Al) having a thickness of 3.0 mm by wave height analysis using the MCA4. The radiation spectrum due to the insertion of the filter is measured.

  As shown in the block diagram of FIG. 3, the high-speed current / voltage amplifier 2 of this embodiment has a configuration using a high-speed operational amplifier (single power supply voltage feedback amplifier: AD8032, trade name of Analog Devices) 25, A 10 kΩ resistor (first resistor) connected to the MPPC 11 of the MPPC radiation detection means 1 and loaded with a 70.7 V bias voltage (first bias voltage) 26 immediately before Geiger mode, and a 5 V bias lower than the first bias voltage 26 A high-speed operational amplifier 26 loaded with a voltage (second bias voltage), a 1 μF microcapacitor and a 100Ω resistor (second resistance) 22 connected between the MPPC 11 of the LSO-MPPC radiation detection means 1 and the high-speed operational amplifier 25, Inversion current consisting of a 51kΩ resistor (third resistor) 24 that bypasses the high-speed operational amplifier 25 A voltage amplification circuit is included.

  The AD8032 (dual) single supply voltage feedback amplifier has high speed performance with a small signal bandwidth of 80 MHz, a slew rate of 30 V / μs, and a settling time of 125 ns. This performance can be obtained by consuming less than 4.0 mW of power from a single +5 V power supply. As a result, the operation time of a high-speed battery drive system or the like can be extended without impairing dynamic performance.

  The MPPC 11 operates immediately before the Geiger mode at the first bias voltage 26 immediately before the Geiger mode and a temperature of 23 ° C. When the pulsed photocurrent 10a generated from the LSO-MPPC radiation detection means 1 including the LSO single crystal 10 and the MPPC 11 flows, a negative voltage is generated on the left side of the second resistor 22 in the figure, and this is provided with the high-speed operational amplifier 25. It is amplified by an inverting current / voltage amplifier circuit and becomes a BNC output 28 to the MCA 4 (FIG. 1) via a coaxial cable.

An event pulse is generated from the BNC output 28, and the generated voltage Vb is expressed as in Equation 1, where Ib is a photocurrent.
[Formula 1]
Vb = 10 × 10 ³ × Ib [V] (1)

  As described above, in the spectrum measurement of radiation (X-ray, γ-ray) using the LSO-MPPC radiation detection means 1, it is important to reduce the dark count rate as much as possible. For this purpose, although the sensitivity of the MPPC 11 is lowered, it is necessary to reduce the dark count rate by making the bias voltage 26 as low as possible. In particular, the event pulse peak value of an X-ray photon having a low photon energy is almost the same as that of a dark count, so it is necessary to measure the bias voltage (Vg) 26 immediately before the Geiger mode. Further, since the bias voltage (Vg) 26 varies depending on the MPPC, it is necessary to measure the bias voltage (Vg) 26 of each MPPC (see FIGS. 4A and 4B). In this example, the bias voltage (Vg) 26 was set to 70.7 V at which the dark count rate was approximately 0 cps.

  In order to compare with the spectrum using the LSO-MPPC radiation detection means 1, first, an X-ray spectrum was measured using a CdTe detector (FIG. 5). As the X-ray tube voltage V increased, the maximum photon energy and the braking X-ray peak energy increased. On the other hand, insertion of a 3 mm thick Al filter absorbed low energy X-ray photons and increased peak energy.

  The CdTe detector uses a high quality CdTe single crystal. This CdTe can be used as an X-ray or γ-ray detector at room temperature, and has recently been attracting attention as a compound semiconductor whose detection efficiency is dramatically improved compared to Si. Applied.

FIGS. 6 (a), 6 (b), and 7 show the results of measuring γ-rays generated from americium 241 ( ²⁴¹ Am) using a CdTe detector and LSO-MPPC radiation detection means 1, respectively. When the CdTe detector of FIG. 6A was used, a sharp γ-ray spectrum of 59.5 keV could be observed. Data of spectrum measurement by event pulse wave height analysis of γ-rays in FIG. 6B is obtained as a change in the number of γ-ray photons depending on the ch number. In the graph of FIG. 6A, the horizontal axis is converted into photon energy. In spectrum measurement, the channel number and energy of γ-ray photons are proportional.

  On the other hand, when the LSO-MPPC radiation detection means 1 was used, a peak of γ rays could be observed, but the energy resolution was 59.5 keV and about 80% (100 × 230/280).

As a result of measuring the spectrum using the LSO-MPPC radiation detection means 1, the number of peak channels and the maximum number of channels increased as the tube voltage V increased as shown in FIG. On the other hand, the insertion of the Al filter did not change the maximum number of channels, and the number of peak channels increased.
(Example 2)

  As shown in FIG. 8, the dark countless radiation detection energy discrimination imaging system (hereinafter abbreviated as high-speed X-ray photon counting CT or PC-CT system) 40 according to the second embodiment of the present invention uses a controller 7 to detect the subject Tg. X-ray tube unit 6 that selectively sets according to characteristics and emits X-rays at a low dose rate, LSO-MPPC radiation detection means 1, high-speed current / voltage amplifier 2, and preset rotation A scanning stage 8 that linearly scans the subject Tg on the LSO-MPPC radiation detection means 1 that is arranged between the subject Tg mounted on the turntable Tt that rotates in steps and the high-speed current / voltage amplifier 2 and attached to the center. And a high-speed comparator that sets the lower limit peak value of the event pulse from the high-speed current / voltage amplifier 2 3 and a counter card 9 that detects an event pulse generated from the high-speed comparator 3 and having a voltage lower than the lower limit voltage Vl and counts square wave pulses, and CT data from the projection data obtained by counting the square wave pulses by the counter card 9 PC 5 for reconstructing a tomographic image.

  In the high-speed X-ray photon counting CT system (PC-CT system) 40, a high-speed comparator 3 and a counter card 9 are added in place of the MCA 4 in the X-ray photon counting and spectrum measurement system 30. A controller 4, a subject Tg, and a turntable (Sigma SP Co., Ltd. SGSP-60YAW-OB) Tt, etc., which are mounted with the subject Tg and rotating at a preset rotation step of 1 to 3 °, are added.

  In this embodiment, the distance between the X-ray tube unit 6 and the LSO-MPPC radiation detection means 1 is l. The distance from the tip of the X-ray tube in the X-ray tube unit 6 to the center of the subject Tg is 960 mm, and the distance from the center of the turntable Tt to the LSO-MPPC radiation detection means 1 is an enlargement ratio of the subject Tg. In order to make it small, it is set to 40 mm.

  The event pulse generated from the high-speed current / voltage amplifier 2 is input to the high-speed comparator 3, and a lower limit peak value (lower limit voltage) Vl that is the lower limit of the event pulse peak value is set. Since the event pulse peak value is proportional to the photon energy, the average energy of X-rays used for CT imaging increases as the lower limit voltage Vl increases. In the PC-CT system 40, the maximum energy corresponds to the tube voltage V. For example, the maximum energy when the tube voltage V is 80 kV is 80 keV. In this embodiment, this PC-CT system 40 is constructed in order to confirm whether or not the energy discrimination effect appears in the CT image due to the change in the lower limit voltage Vl.

X-ray photons Xr that pass through the subject Tg are amplified using the LSO-MPPC radiation detection means 1 and the high-speed current / voltage amplifier 2. Next, the lower limit voltage Vl of the event pulse is set by the high speed comparator 3, the event pulse equal to or higher than the lower limit voltage Vl is detected, and the square wave pulse generated from the high speed comparator 3 is counted by the counter card 9. In the PC-CT system 40, projection data of the subject Tg is acquired by linear scanning of the subject Tg and the LSO-MPPC radiation detection means 1, and the subject Tg is rotated by 1 to 3 degrees of a preset rotation step. A CT tomographic image is reconstructed from the projection data obtained by repeatedly controlling the linear scan and rotation of the subject Tg by the biaxial controller 4.
(X-ray tomography)

  Here, using the PC-CT system 40, X-ray tomography was performed under the conditions of an X-ray tube voltage V = 80 kv, a scan step of 0.5 mm, and a rotation step of 1.0 °. The maximum and minimum relative counts in the reconstructed CT image pixels are shown as black and white, respectively. On the other hand, the maximum value and the minimum value of density in the obtained JPEG image file are represented by white and black, respectively.

  First, CT tomographic images obtained by the PC-CT system 40 of iodine contrast agents in two polyethylene vials containing iodine contrast agents having concentrations of 15 mg / ml and 30 mg / ml, respectively, are shown in FIGS. Show. When the X-ray tube voltage V = 80 kV was constant, the image density difference of the iodine contrast agent decreased as the event pulse lower limit voltage Vl by the high-speed comparator 3 increased. This is because the average photon energy of CT imaging increases with an increase in the lower limit voltage Vl.

  In addition, CT tomographic images of an acrylic phantom with two holes are shown in FIGS. Similar to FIGS. 9A to 9D, iodine contrast agents having different concentrations are contained, and the image density difference of the iodine contrast agent is reduced as the event pulse lower limit voltage Vl is increased by the high-speed comparator 3.

Next, CT tomographic images of the dog's heart are shown in FIGS. The coronary artery of the dog's heart is injected with iodine microspheres having a diameter of 15 μm. As shown in FIGS. 11A to 11D, the contrast between the myocardium and the coronary artery decreased as the event pulse lower limit voltage Vl increased by the high-speed comparator 3.
(Example 3)

As described above, the X-ray photon counting and spectrum measurement system 30 and the high-speed X-ray photon counting CT (PC-CT) system 40 using the LSO-MPPC radiation detection means 1 have been described. The YAP (Ce) -MPPC detector Although the details are omitted, the same spectrum and CT imaging results as those of the X-ray photon counting and spectrum measurement system 30 and the PC-CT system 40 were obtained.
Example 4

Also, when using a gadolinium contrast agent, although details are omitted, a change in contrast of the gadolinium contrast agent was observed due to a change in the event pulse lower limit voltage Vl, as in Example 2. This is because the average photon energy for CT imaging increases with an increase in the event pulse lower limit voltage Vl.
(Key points obtained from the examples)

  In the first and second embodiments, the MPPC 11 having a pixel pitch of 50 × 50 μm (400 pixels) was used, but experiments using 25 × 25 μm (1600 pixels) and 100 × 100 μm (100 pixels) were performed. I have not been told. However, since the dynamic range is wide at 1600 pixels, it is considered that the energy resolution is also improved.

  Since the energy resolution of the LSO-MPPC radiation detection means 1 is low, K-edge imaging using an iodine or gadolinium contrast agent is difficult, but dual energy subtraction is possible.

  Further, in the X-ray photon counting, a strict electromagnetic shield is required in the CdTe detector, but since the peak photocurrent value of MPPC is larger than that of the photodiode, there is no need for the electromagnetic shield. The LSO-MPPC radiation detection means 1 is a great advantage in the construction of a medical image system such as a CT system.

  The results obtained from the above examples are summarized as follows.

1. Using the LSO-MPPC radiation detection means 1, the functions of both the X-ray photon counting and spectrum measurement system 30 and the PC-CT system 40 were confirmed. The dark count rate was an extremely low value of 0.5 cps or less at the MPPC bias voltage Vg = 70.7V. ^In monochromatic γ-ray photon counting using a ²⁴¹ Am source and MCA, the energy resolution of the LSO-MPPC radiation detection means 1 was 59.5 keV and almost 100%. However, the peak energy of bremsstrahlung photons is well comparable to the value obtained using a CdTe detector, and the energy resolution improved with increasing gamma photon energy.

2. The width of the event pulse is about 200 ns, and the event pulse width needs to increase beyond 5 μs in order to correctly detect the event pulse height. The event pulse height H (V) is expressed as Equation 2 as a function of the number of channels N (channel) determined using MCA4.
[Formula 2]
H = 2.5 × 10 ⁻³ N (2)
However, the event pulse height measured using MCA4 decreased as the event pulse width decreased due to the sampling period of the AD converter in MCA4.

  3. When using the LSO-MPPC radiation detection means 1, l. Since it was difficult to amplify the X-ray photocurrent from the MPPC using a 4 MHz band gain operational amplifier LMC662 or a commercially available charge amplifier, a high speed current / voltage amplifier 2 having a high speed operational amplifier (AD8032) 25 of 80 MHz band is used. developed. In an initial experiment, when a non-inverting current / voltage amplifier circuit was used using a 7 ns delay comparator AD8561, a false oscillation from the comparator was observed, and the count rate was greatly increased by the oscillation. In order to prevent oscillation, an inversion current / voltage amplification circuit with a hysteresis circuit was developed. Subsequently, a dual energy comparator means for determining the energy range (both level and width) is desired.

  4). In the PC-CT system 40, the CT tomographic image density difference between the two iodine contrast agents decreased as the event pulse lower limit voltage Vl increased due to the increase in the average photon energy of CT. However, the CT tomographic image contrast at Vl = 0.5V was almost equal to that obtained using IK edge image processing with a CdTe detector. In the second embodiment, the X-ray tube voltage V is selected to be 80 kV in order to observe the variation of the CT tomographic image contrast due to the change in the lower limit voltage Vl. However, since the high energy X-ray photon reduces the image contrast, the image contrast is reduced. In order to improve, it is desirable to select the X tube voltage V in the range of 60 to -70 kV. Because of photon counting by the scintillation method, it is difficult to perform K-edge image processing using an iodine contrast agent that requires high energy resolution. However, it is possible to perform dual energy subtraction using dual energy comparator means.

  5. ED (Energy Dispersive)-in basic experiments on CT, A CdTe detector with an energy resolution of 2 keV is usually used. On the other hand, the LSO-MPPC radiation detection means 1 provided with the high-speed current / voltage amplifier 2 can easily increase the count rate beyond lMcps and realize a PC-CT having an ED (energy discrimination) function. Can do.

  6). Since the maximum count rate of the MCA 4 used in Example 1 above for measuring the X-ray spectrum is 10 kcps, the dark count rate of the LSO-MPPC radiation detection means 1 is reduced by reducing the bias voltage of the MPPC 11. Minimized. The MPPC gain decreases as the bias voltage decreases, but the brightness of LSO with a light emission lifetime of 40 ns is 76% higher than that of thallium activated sodium iodide [Nal (TI)] with a light emission lifetime of 1 μs, which is considerably large. An event pulse with a wave height is generated.

7). The graininess of the CT tomographic image increases due to a decrease in the number of photons per measurement point when the event pulse lower limit voltage Vl increases. When the scattering of the X-ray photon Xr from the subject Tg is not taken into consideration, in order to improve the graininess at a constant X-ray tube voltage V, it is desirable to increase the X-ray tube current beyond 1.5 mA. The spatial resolution is basically determined by the X-ray photocathode size, and the photocathode size in the spatial resolution of Example 1 is l. It was 0 × 1.0 mm ² . In order to improve the spatial resolution, it is desirable to reduce the size of the LSO crystal. In addition, CT tomographic image quality is improved by reducing CT scanning and rotation steps.

  8). In the PC-CT system 40 of the present invention, the X-ray tube voltage V = 80 kV, the X-ray tube current = 1.5 mA, the event pulse lower limit voltage Vl = 0.5 V, the maximum count rate is 250 kcps, and per measurement point. The number of photons was calculated to be 5 kc (= 250 kc / 50). Furthermore, by decreasing both the event pulse lower limit voltage Vl and increasing both the X-ray tube voltage V and current, the count rate increased significantly beyond 1 Mcps.

  9. From the above, the LSO-MPPC radiation detection means 1 considers the energy resolution and the maximum count rate, and performs ED (energy discrimination) for performing CT imaging and dual energy subtraction image processing similar to the K edge enhancement method. It can be applied to a next generation PC-CT system having a function. Furthermore, the LSO-MPPC radiation detection means 1 can also be applied to both single photon emission computed tomography (SPECT) and PET using an LSO crystal having a large specific gravity. Thus, since the LSO-MPPC radiation detection means 1 can easily discriminate high-energy γ-ray photons, it is possible to realize energy discrimination type PET / CT and SPECT / CT.

  The foregoing descriptions of specific embodiments of the present invention have been presented for purposes of illustration. They are not intended to be exhaustive or to limit the invention to the precise form described. It will be apparent to those skilled in the art that many modifications and variations are possible in light of the above description.

  The present invention uses an LSO-MPPC radiation detection means 1 that detects radiation such as X-rays and γ-rays, which is a combination of an LSO single crystal, which is a scintillator with a short emission lifetime, and MPPC, and performs imaging by counting or converting X-rays. Performs energy discrimination imaging by selecting photon energy or average energy, has a simple structure and is excellent in cost and X-ray to electric signal conversion efficiency, and can reduce the amount of X-ray exposure to the subject. Since the dark countless radiation detection energy discriminating imaging system 40 can be provided, a diagnosis of a soft tissue disease from the orthopedic or oral surgery field, which diagnoses and treats a hard tissue disease such as a bone or a tooth by simple X-ray examination A wide range of medical fields covering internal medicine and surgical fields, and other X-ray images It can be universally applied to industries related to.

1 LSO-MPPC radiation detection means (LSO-MPPC detector)
2 High-speed current / voltage amplifier 3 High-speed comparator 4 Multi-channel analyzer (MCA)
5 Computer unit (PC)
6 X-ray tube unit or X-ray tube 7 Controller 8 Scan stage 9 Counter card 10 LSO or LSO single crystal 10a Pulse photocurrent 11 MPPC (multi-pixel photon counter or multi-pixel APD)
12 Ceramic substrate 13 Black rubber 14 Electrode 15 BNC connector 16 Aluminum (Al) cap 16a Thin film window (photoelectric surface)
16b Al case 21 Capacitor 22, 23, 24 Resistor 25 High-speed operational amplifier (single power supply voltage feedback amplifier)
26, 27 Bias voltage 28 BNC output 30 X-ray photon counting and spectrum measurement system 40 Dark countless radiation detection energy discrimination imaging system (High-speed X-ray photon counting CT system: PC-CT system)
Tg Subject Tt Turntable V X-ray tube voltage Vl Event pulse lower limit voltage (Lower limit crest value)
Xr X-ray or X-ray beam or X-ray photon

Claims (6)

An X-ray tube unit that is selectively set by the controller according to the characteristics of the subject and emits X-rays at a low dose rate;
It is enclosed in an aluminum (Al) cap having a thin film window that becomes an X-ray photocathode, and is laminated on a multi-pixel photon counter (hereinafter referred to as MPPC) formed on a ceramic substrate and connected to a pair of electrodes. LSO-MPPC radiation (X-rays, γ-rays) detecting means for receiving X-rays radiated from the X-ray tube unit and transmitted through the subject, comprising Lu2 (SiO4) O (hereinafter referred to as LSO) single crystal; ,
A high-speed current / voltage amplifier that amplifies the photocurrent generated from the LSO-MPPC radiation detection means and generates an event pulse;
A multi-channel analyzer (hereinafter referred to as MCA) for analyzing the pulse height of event pulses from the high-speed current / voltage amplifier,
And an X-ray photon counting and spectrum measuring system comprising: a computer unit (hereinafter referred to as PC) that represents a pulse height analyzed event pulse from the MCA as a change in a photon number channel (hereinafter referred to as ch). Dark countless radiation detection energy discrimination imaging system. An X-ray tube unit that is selectively set by the controller according to the characteristics of the subject and emits X-rays at a low dose rate;
The X-ray tube unit comprising an LSO single crystal enclosed in an Al cap having a thin film window serving as an X-ray photocathode and formed on a ceramic substrate and laminated on an MPPC to which a pair of electrodes are connected LSO-MPPC radiation detection means for receiving X-rays emitted from and transmitted through the subject;
A high-speed current / voltage amplifier that amplifies the photocurrent generated from the LSO-MPPC radiation detection means and generates an event pulse;
A high-speed comparator that sets a lower limit peak value of the event pulse from the high-speed current / voltage amplifier;
A counter card that detects event pulses above the lower limit voltage generated from the high-speed comparator, and counts square wave pulses;
A dark countless radiation detection energy discrimination imaging system, comprising: a PC for reconstructing a tomographic image from projection data obtained by counting square wave pulses with the counter card. An X-ray tube unit that is selectively set by the controller according to the characteristics of the subject and emits X-rays at a low dose rate;
The X-ray tube unit comprising an LSO single crystal enclosed in an Al cap having a thin film window serving as an X-ray photocathode and formed on a ceramic substrate and laminated on an MPPC to which a pair of electrodes are connected LSO-MPPC radiation detection means for receiving X-rays emitted from and transmitted through the subject;
A high-speed current / voltage amplifier that amplifies the photocurrent generated from the LSO-MPPC radiation detection means and generates an event pulse;
A linear scan of the subject is performed by the LSO-MPPC radiation detection means disposed between the subject mounted on a turntable that rotates at a preset rotation step and the high-speed current / voltage amplifier and provided at the center. A scanning stage to
A high-speed comparator that sets a lower limit peak value of the event pulse from the high-speed current / voltage amplifier;
A counter card that detects event pulses above the lower limit voltage generated from the high-speed comparator, and counts square wave pulses;
A dark countless radiation detection comprising: a PC for reconstructing a tomographic image from projection data obtained through the counter card by repeatedly controlling linear scanning and rotation of the subject with a two-axis controller Energy discrimination imaging system. The LSO-MPPC radiation detection means includes:
An Al cap having a thin film window that shields and shields light and serves as an X-ray photocathode;
A laminate of the LSO single crystal and MPPC sealed in the Al cap and affixed to the photocathode;
And a BNC connector for accommodating the Al cap therein,
Receiving X-rays transmitted through the subject through the thin film window, detecting (counting) X-ray photons from the front surface of the LSO single crystal;
4. The X-ray photon that has passed through the MPPC is Compton scattered by the ceramic substrate behind it to generate scattered radiation and be converted into fluorescent radiation. Dark countless radiation detection energy discrimination imaging system.   4. The X-ray tube unit generates X-rays at a low dose rate by reducing an X-tube current in a range of 1.0 μA to 2.0 mA order. 5. The dark countless radiation detection energy discrimination imaging system according to claim 1. The high-speed current / voltage amplifier is:
A first resistor coupled to the MPPC of the LSO-MPPC radiation detection means and loaded with a first bias voltage immediately before Geiger mode;
A high-speed operational amplifier (single power supply voltage feedback amplifier) loaded with a second bias voltage lower than the first bias voltage;
A micro capacitor and a second resistor connected between the MPPC and the high-speed operational amplifier;
The dark countless radiation detection energy discrimination imaging system according to any one of claims 1 to 3, further comprising an inverting current / voltage amplification circuit including a third resistor that bypasses the high-speed operational amplifier. .

Images