JP2010243395A - X ray-gamma ray imaging device - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide an X ray-gamma ray imaging device using X rays and gamma rays capable of reliably obtaining functional information by gamma rays and form information by X rays without alignment. <P>SOLUTION: The X ray-gamma ray imaging device includes a detector 13 on which pixels are two-dimensionally arranged. The detector 13 includes an output structure which includes a photon counting type radiation detection circuit having two or more discriminators for radiation energy for each pixel and independently outputs signal detection information of each pixel and energy affiliation information by the discriminators on the basis of the unit of pixel. A collimator 14 having a cone beam shape for directing only the X ray beam from an X ray tube 12 disposed opposite the detector 13 is disposed at the front surface of the detector 13, and the collimator 14 simultaneously collimates the gamma rays radiated from an object in the field of view to enter the detector 13. <P>COPYRIGHT: (C)2011,JPO&INPIT

Description

  The present invention relates to an X-ray / gamma ray imaging apparatus that captures an image inside a subject using both X-rays and gamma rays.

In the medical field, an energy signal integrating detector that measures the fluence of incident X-rays is generally used as a two-dimensional X-ray detector. This is because, in the case of X-ray, very much, apply the photon counting technique for counting individual photons are used such as in nuclear medicine for respective energy photon number 1M~2Gcps / mm ² incident on the detector This is because it is difficult to do. For example, in a scintillation camera used in the field of nuclear medicine diagnosis, the number of photons of gamma rays is about 200 to 300 kcps per detector, and the count is 1 / 10,000 of the counting ability required for medical X-ray detection. It is less than ability. As a result, it is extremely difficult to use the photon counting technique in medical X-ray flat panels and X-ray CT scanners, and there is no detector capable of simultaneously counting X-rays and gamma rays.

In addition, the photon detection technology used in SPECT (Single Photon Emission Computed Tomography), which visualizes the function of internal organs, uses a scintillator such as sodium iodide and a photomultiplier tube. In the photon detection method, the intrinsic spatial resolution of the detector is lowered due to the diffusion of light, fluctuation of the photon energy, and the position of the photomultiplier tube. As a result, the spatial resolution of the SPECT image is 5 to 8 mm for clinical use. In addition to making it impossible to detect a radiopharmaceutical in a small lesion, it is very difficult to identify an accumulation position from a SPECT image alone. Accurate quantification is not possible, and this makes it difficult to use for treatment planning and treatment effect determination.

  In order to compensate for such a low spatial resolution of the SPECT image, an X-ray CT apparatus is installed in addition to the SPECT apparatus, and a single image in which functional information by SPECT and morphological information by X-ray CT are superimposed can be created. Has been done. However, these devices simply have two types of modalities adjacent to each other, and the alignment of images reconstructed with different geometries is complicated. Further, in order to improve the quantitative property of a SPECT image by using such an SPECT / CT integrated apparatus, accurate attenuation correction of gamma rays is necessary, but a signal that integrates the entire energy of the current incident X-rays. Even if X-ray CT images obtained by the processing method are used, there is a problem that an accurate line attenuation coefficient cannot be obtained, and a true quantitative SPECT image can be obtained simply by using a conventional X-ray CT apparatus. The current situation is not possible.

  In recent years, monoclonal antibodies are labeled with imaging nuclides (gamma-emitting nuclides) and therapeutic nuclides (beta-emitting nuclides), respectively, and the size and spread of tumors are accurately grasped with the first labeled antibodies. After confirming that accumulation in a normal organ has not been performed, an isotope internal therapy in which an antibody labeled with a therapeutic nuclide is administered has been used. This is a technique that has received the most attention in the future cancer treatment because it is minimally invasive and can specifically destroy a specific cancer at the cellular level. Since such imaging nuclides cannot be realized with positron emitting nuclides, it is impossible to use PET (Positron Emission Tomography), and it is necessary to rely on SPECT technology. On the other hand, as described above, the SPECT image has a low spatial resolution in the image by the current photon detection method, and even if an attempt is made to quantify with the gamma ray attenuation correction using the X-ray CT, Since the photon measurement method cannot obtain an accurate distribution of the linear attenuation coefficient, the exact position of the monoclonal antibody cannot be identified, which is not sufficient for the treatment plan and the determination of the treatment effect.

  The present invention has been made in view of the above circumstances, and in particular, X-rays and gamma rays can be reliably obtained without aligning functional information by gamma rays and morphological information by X-rays with respect to the same subject. It is an object of the present invention to provide a new modality to be called an X-ray / gamma-ray imaging apparatus using the.

  In order to achieve the above object, an X-ray / gamma-ray imaging apparatus according to the present invention includes a photon counting radiation detection circuit having two or more radiation energy discriminators in a radiation detector in which pixels are two-dimensionally arranged. And an output structure for independently outputting signal detection information of each pixel and energy affiliation information by a discriminator in units of pixels, and an X-ray from an X-ray tube installed facing this detector. Basically, a cone beam collimator for directing only to a line beam is arranged in front of the detector, and the collimator collimates gamma rays emitted from an object in the field of view and incident on the detector at the same time. Features.

  As described above, it is possible to solve the disadvantages of a single imaging apparatus using X-rays and gamma rays, and to provide an unprecedented epoch-making image and usage that can realize the quantitativeness of image measurement.

FIG. 1 is a diagram showing an outline of the configuration of an X-ray / gamma ray imaging apparatus according to an embodiment of the present invention. FIG. 2 is a view for explaining the geometry of the arrangement of an X-ray tube and a detector showing a modification of the embodiment. FIG. 3 is a graph showing an example of energy spectra of X-rays and gamma rays. FIG. 4 is a graph showing another example of energy spectra of X-rays and gamma rays. FIG. 5 is a diagram showing an example of a data collection sequence suitable for the case where the energy spectra of X-rays and gamma rays are shown in FIG. FIG. 6 is a diagram showing an example of a data collection sequence suitable for the case where the energy spectra of X-rays and gamma rays are shown in FIG. FIG. 7 is a block diagram showing a schematic configuration of a detector mounted on the X-ray / gamma ray imaging apparatus according to the embodiment.

  An embodiment of an X-ray / gamma ray imaging apparatus according to the present invention will be described with reference to the accompanying drawings.

  FIG. 1 shows an outline of an X-ray / gamma ray imaging apparatus according to this embodiment. This X-ray / gamma-ray imaging apparatus is a composite imaging apparatus in which an X-ray CT scanner using X-rays and a SPECT apparatus using a gamma camera are integrated. Here, the term “integration” does not mean that both modalities are simply placed on one stand, but “integration of X-ray and gamma-ray collection systems and processing systems for collected data of these different types of radiation”. It is integrated (including some integrations) or in a state where it can be used mutually. "

  As shown in FIG. 1, the X-ray / gamma-ray imaging apparatus includes a gantry 1 so that a top plate 2A of a bed 2 is loosely inserted in an imaging space S formed in a state of penetrating the gantry 1 so as to advance and retreat. It has become. The subject P is placed on the top board 2A, and a tomographic image of the subject P is taken and / or a SPECT image is taken.

  The X-ray / gamma-ray imaging apparatus further includes an X-ray tube 12 (X-ray source) and a detector 13 that are installed inside the gantry 1 so as to face each other by a support 11. When the support 11 rotates, the pair of the X-ray tube 12 and the detector 13 can rotate around the body axis direction of the subject P while maintaining the opposing geometric relationship. A cone beam collimator 14 having a geometrical shape that converges at the X-ray exposure position of the X-ray tube is disposed on the radiation (X-ray, gamma ray) incident surface of the detector 13.

  In addition, the X-ray / gamma ray imaging apparatus includes an X-ray tube high voltage generator 15, a bed driver 16, a detection system driver 17, and a data collector 18 installed on the gantry 1. Further, the X-ray / gamma-ray imaging apparatus includes an X-ray controller 21 outside the gantry 1, a computer 22 that operates under program calculation by the CPU, and a console 23 (an interface between the computer 22 and the operator) ( An input device 24 and a display device 25) and an image storage device 26.

  Among the components described above, the X-ray tube 12 emits X-rays in response to a pulsed or continuous wave high voltage signal supplied from the high voltage generator 13. The high voltage generator 13 generates an X-ray control signal based on a predetermined pulse sequence from the X-ray controller 21 and sends this signal to the high voltage generator 13. For this reason, since the high voltage generator 13 supplies a high voltage signal to the X-ray tube 2 according to the X-ray control signal, the X-ray exposure described above can be performed. The computer 22 activates the pulse sequence for X-ray exposure and sends the sequence information to the X-ray controller 21, whereby the X-ray controller 21 operates.

  The gantry driver 16 operates in response to a control signal given from the computer 22, and as a result, the top plate 2 </ b> A of the gantry 2 can be moved up and down or in the longitudinal direction thereof. The detection system driver 17 operates in response to a control signal given from the computer 22 and can drive the support 11 to rotate.

  The data collector 18 collects data corresponding to the number of photons of X-rays and gamma rays detected by the detector 13 and sends this data to the computer 22.

  The computer 22 includes components for operating as a computer, such as a memory, in addition to the CPU. The CPU activates a program stored in the memory in advance, and X-rays collected via the data collector 18 and A tomogram is reconstructed or created based on data corresponding to the number of photons of gamma rays, and the operations of the X-ray controller 21, the gantry driver 16, and the detection system driver 17 are controlled. Various programs for this purpose are stored in advance in the memory of the computer 22 as an example. On the display 25 of the console 23, the tomogram reconstructed or created and necessary information are displayed under the control of the computer 22. The input device 24 is used by the operator for the operator to input operational instructions and subject information. The image storage device 26 functions as an external storage unit of the computer 22 and stores data of a reconstructed or created image.

  In the case of the apparatus configuration shown in FIG. 1, the number of pairs of the X-ray tube 12 and the detector 13 is one, but this is not necessarily limited. For example, as shown in FIG. 12B, 12C) and three pairs of detectors 13A (13B, 13C) may be arranged around the imaging space S of the gantry 1 at equal angular intervals. Thereby, the amount of movement of the X-ray tube 12 and the detector 13 can be reduced, and data can be collected in a short time.

  Of the components of the X-ray / gamma-ray imaging apparatus of FIG. 1 described above, the detector 13 will be described in more detail.

The detector 13 makes X-rays that have passed through the subject P incident and outputs an electrical signal in response to the number of incident photons. On the other hand, this detector 13 counts single photons (single photons) emitted from gamma-ray emitting nuclides such as ^99m Tc, ²⁰¹ Tl, ¹²³ I, and ¹³³ Xe that are administered to the subject. The electric signal according to is output.

  That is, the detector 13 is configured as a radiation detector in which pixels are two-dimensionally arranged. This detector 13 has a photon counting radiation detection circuit having two or more radiation energy discriminators TH1, TH2,... (See FIGS. 4 and 6) for each pixel, and signal detection information of each pixel. And an output structure for independently outputting energy affiliation information by the discriminator in units of pixels.

  Here, with reference to FIG. 7, a schematic electrical configuration of the detector 13 will be described as a block diagram.

  The detector 13 includes four elements and an electric circuit including a scintillator 121 that receives X-rays and gamma rays, a FOS (Fiber Optics Plate) 122, a CMOS photoelectric converter 123, and a processing circuit 124 including a CMOS preamplifier. It is formed in a plate shape or a layer shape and has a structure in which they are stacked on each other. A block diagram from the scintillator 121 to the processing circuit 124 is shown in FIG.

  The scintillator 121 is a plate formed of a CsI material that converts radiation into light, and has an incident surface on which radiation is incident and an exit surface that emits light in a back-to-back state. The FOS 122 is a bundle of optical fibers having a small diameter (5 to 10 μm in diameter) bundled and spread upright in a plane shape. The bundle is arranged so that the axial direction of each optical fiber is the incident direction of radiation. The scintillator 121 and the photoelectric converter 123 are interposed.

  The photoelectric converter 123 is a semiconductor layer that converts light incident through the FOS 22 into an electrical signal, and a photoelectric conversion element is formed in each pixel by CMOS on the substrate. That is, the size of each photoelectric conversion element to be formed corresponds to a pixel, and the size of the pixel is, for example, 200 μm × 200 μm. Note that the vertical and horizontal sizes of each pixel may be different from each other.

Further, the processing circuit 124 is also formed of CMOS on the substrate, and amplifies the electric signal for each pixel output from the photoelectric converter 123. The electrical signals for each pixel are then input in parallel to the comparison input terminals of the six types of comparators 125 _{1 to} 125 ₆ as shown in FIG. These the reference input of the comparator ₁₂₅ to 253 _6, different thresholds for the energy discrimination DC1～DC6 (e.g. DC1 <DC2, ..., DC5 < DC6) is given. For this reason, six discriminators are formed in accordance with the energy level of the X-ray photons input to a certain pixel. The output of the comparator ₁₂₅ 1 to 125 ₆ are respectively connected to a counter ₁₂₆ 1 to 126 _6, the count value of their counter ₁₂₆ 1-126 ₆ is read out in serial format. Therefore, the energy of X-rays (photons) for each pixel is discriminated by the discriminator, and the number of X-ray photons having energy belonging to each energy region is measured by the counter 126 ₁ (˜126 ₆ ) in the next stage, It is output serially as a measured value of digital quantity.

  As described above, the collimator 14 is disposed on the front surface of the detector 13, that is, the radiation incident surface. The collimator 14 is configured as a cone beam-shaped collimator that directs only to the X-ray beam from the X-ray tube 12 installed facing the detector 13. As a result, not only X-rays but also gamma rays radiated from the object in the field of view of the detector 13, that is, the collimator 14 and incident on the detector 13 can be collimated in parallel (simultaneously).

  Here, typical examples of energy spectra of X-rays and gamma-ray emitting nuclides to be used are shown in FIGS. In these figures, the horizontal axis represents the energy of radiation, and the vertical axis represents the number of photons (count value). As can be seen from these, the number of X-ray photons is overwhelmingly larger than that of gamma rays. As the next trend, as shown in FIG. 3, it can be said that the spectral distributions of X-rays and gamma rays overlap each other on the axis representing energy, and as shown in FIG. It can be said that they are separated. In the case of the present embodiment, as shown in FIG. 3, when “the spectral distributions of X-rays and gamma rays overlap each other”, the acquisition sequence shown in FIG. 5 is used. Further, as shown in FIG. 4, when “the spectral distributions of X-rays and gamma rays do not overlap each other”, the acquisition sequence shown in FIG. 6 is used.

  In the case of the acquisition sequence of FIG. 5, X-ray exposure and gamma ray acquisition are alternately performed in a time division manner. X-ray exposure may be performed by supplying a pulsed high voltage to the X-ray tube 12, and when passing through the subject P, X-ray photons are almost regarded as particles and detected by the detector 13. On the other hand, since gamma rays are always emitted from the gamma-ray emitting nuclide administered to the subject P, the photons of the gamma rays are also regarded as particles by the detector 13 and detected by the detector 13. However, since the number of X-ray photons is overwhelmingly larger than that of gamma rays, even if the photons of gamma rays enter the detector 13 during the collection of X-ray photons, the effect can be ignored. For this reason, the X-ray exposure period Tx is set as a pulse high voltage application period to the X-ray tube, and when this X-ray exposure period Tx ends, a certain period of gamma ray acquisition period Tγ is set. The detector 13 continuously collects without distinguishing between the two periods Tx and Tγ. Since both the periods Tx and Tγ are identified by the computer 22, the data collected in the respective periods Tx and Tγ can be recognized as information indicating the number of photons of X-rays and gamma rays.

  On the contrary, in the case of the acquisition sequence of FIG. 6, X-ray exposure and gamma ray acquisition are executed in parallel. Gamma rays are also collected at the time of X-ray collection, and gamma ray photons may also enter the X-ray photon collection data as noise, but the amount can be ignored, so parallel collection can be performed in this way. It is. In the case of the radiant energy discriminators TH1, TH2,... Shown in FIG. 6, the energy spectra of the X-rays and the gamma rays are separated. For example, the lower discriminators TH1, TH2, TH3, TH4. Thus, discrimination of X-ray energy can be performed for each pixel, and discrimination of gamma-ray energy can be performed for each pixel by the upper discriminators TH4, TH5, TH6. The number and value of the discriminators TH1, TH2,..., That is, the energy window width is arbitrary.

  In this way, as in the collection sequence of FIG. 5, X-rays are intermittently irradiated and detected by the detector 13, and gamma rays are always emitted from the object and detected by the detector 13. Deterioration of the counting characteristics due to almost counting off. In addition, it can be regarded as an almost X-ray transmission image during X-ray imaging, and can be regarded as a radiation image of gamma rays when imaging only gamma rays, and X-rays and gamma rays can be separated and collected by a discriminator.

  In addition, the detector 13 and the X-ray tube 12 are installed opposite to the object, and simultaneous tomography of X-ray CT and SPECT is possible by the rotation data collection mechanism, that is, the support 11, the detection system drive system 17, and the computer 22). become. In this case, the X-ray CT image and the SPECT image can be superimposed and displayed without performing processing such as image enlargement, reduction, or alignment.

  In addition, the target of the SPECT image is an isotope in which a monoclonal antibody is labeled with a gamma-ray emitting nuclide for imaging, its accumulation amount and site are quantitatively grasped from the SPECT image, and the same antibody is labeled with a therapeutic radiation-emitting nuclide. In internal therapy, the treatment plan and treatment effect determination can be performed from the SPECT image and the X-ray CT image.

  In this embodiment, the energy discriminator is set in the same manner as the energy window setting of the nuclide used in SPECT acquisition, and separately from the X-ray CT image obtained from the energy window information set in X-ray data acquisition. The linear attenuation coefficient of the energy of the nuclide can be calculated, and the SPECT image can be quantified.

A direct conversion type semiconductor detection element such as CdTe, CdZnTe, TlBr, or HgI ₂ can be used as a detection element for detecting radiation such as X-rays or gamma rays used in the above-described detector. Of course, instead of such a direct conversion type, a radiation detector combining a scintillator such as CsI and a photoelectric conversion element may be used. Furthermore, it is possible to increase the light collection efficiency and prevent scattered light by interposing an FOS (fiber optic plate) between the scintillator and the photoelectric conversion element.

  In addition, the radiation detection element is a scintillator such as ZnO or GOS processed into a columnar shape, and the columnar scintillator is optically separated and arranged in a two-dimensional shape and a photoelectric conversion element combined. Good. As the photoelectric conversion element, a two-dimensional element such as a CMOS or a CCD can be used. Further, the photoelectric conversion element may be a two-dimensional element composed of a photodiode.

  Furthermore, according to the X-ray / gamma-ray imaging apparatus according to the present embodiment, the X-ray obtained by fixing the positions of the X-ray tube 12 and the detector 13 and relatively moving only the imaging object in the body axis direction. Diagnosis using a line image or a gamma ray planar image obtained by fixing the position of the detector 13 without performing X-ray irradiation and moving only the imaging object in the body axis direction is also possible. . Furthermore, it is also possible to acquire a still image that is captured in a state where the position of the detector 13 is fixed and the imaging object is fixed without moving without performing X-ray irradiation. These various imaging methods are executed under the imaging sequence that the computer 22 has in advance. In the case of this diagnosis, an X-ray image and a gamma-ray planar image can be taken at the same time, and these two types of images can be superposed and diagnosed without performing enlargement / reduction or alignment processing. In place of this overlay diagnosis, a diagnosis combining an X-ray image or a gamma-ray planar image and an X-ray CT image or SPECT image obtained by the present X-ray / gamma ray imaging apparatus is also possible.

  Furthermore, according to the present embodiment, when collecting gamma ray data, a plurality of energy windows can be designated, and the scattered ray correction in the SPECT image can be performed from this data.

  Further, when collecting photons of X-rays and gamma rays, these are collected at the same time, and data collection for scattering correction of a SPECT image is performed using, for example, the TEW method other than at the time of X-ray irradiation: Triple Energy Window method).

  Further, instead of measuring the number of photons of X-rays and gamma rays within the set energy window, it is also possible to use this data by measuring the integrated value of photon energy.

  Furthermore, when collecting X-ray data, multiple energy windows are provided, and these data are used to indicate the effects of beam hardening on X-ray images, contrast enhancement in low-contrast areas, enhancement of calcification, and the extent of the index. It can also be used to calculate

The above is summarized as follows. In the present invention, the detector system is renewed, and the radiation detector is composed of fine pixels (50 μm × 50 μm to 1000 μm × 1000 μm) having a pixel structure, and a preamplifier and two or more discriminators are mounted for each pixel. Each pixel can output energy information and a count value independently, and realize 1 M to ² Gcps / mm ² in the count rate characteristic of each pixel. This realizes a detector that is capable of handling both X-rays and gamma rays with a counting rate characteristic that can withstand medical X-ray measurement despite being a photon counting type. At the same time, a cone-beam collimator focusing on the X-ray generation position is placed in front of the detector to remove scattered rays generated from the object to be imaged. It also has the function of limiting to the gamma rays necessary for imaging of gamma rays from. The size, length, partition wall thickness, and the like of the collimator hole are variously prepared in consideration of the detection sensitivity and resolution of gamma rays to be imaged and the effect of removing scattered X-rays.

  For example, in medical applications, it is possible to collect data for the same cross section in X-ray CT and SPECT using such a detector simultaneously or in a time-sharing manner, and there is no movement of the patient. Positioning in superposition becomes unnecessary. At this time, if the X-ray energy and the SPECT nuclide energy are similar, the X-ray is irradiated in a time-sharing manner, SPECT data collection is performed when the X-ray is not irradiated, and the photon of the photon is irradiated when the X-ray is irradiated. The amount of X-ray data will be collected. This is because the radioactivity of the radiopharmaceutical to be administered is determined, and the number of photons emitted from the radioactivity is about 1 / 10,000 of the number of photons in normal X-ray photography. This is because the effect of gamma rays can be ignored. If the photoelectric peak energies of the X-ray energy region and the gamma ray are different, different types of data can be collected simultaneously by distinguishing the energy windows for collection. Even in such a case, for example, gamma ray data can be collected in an energy window for collecting X-rays other than at the time of X-ray irradiation, and can be used for correcting scattered rays of gamma rays.

  Also, if the energy discriminator is set to be similar for X-rays and gamma rays, the value of the line attenuation coefficient obtained from the X-ray CT image can be used for correction of the SPECT image. Even so, it is also possible to determine the distribution of the line attenuation coefficient by setting multiple energy windows, identifying the medium from the measured X-ray data, and referring to the value of the gamma ray energy for that medium. . Thereby, quantification of a true SPECT image can be realized, and an epoch-making effect is expected also in the area of diagnosis and treatment.

Furthermore, by utilizing such features that X-ray photons can be discriminated and collected for each energy, reduction of beam hardening artifacts generated in X-ray CT images, contrast enhancement of low-contrast areas such as soft tissues, intravascular It is also possible to use for calculation of an index for displaying the degree of calcification such as.

  In addition, X-ray images obtained by a general photon counting method can reduce the amount of irradiated X-rays by several minutes due to the effect of suppressing noise contamination caused by the processing circuit system and the ability to express the X-ray intensity as a count value. It is also possible to reduce the radiation dose to patients. Another feature is that high spatial resolution can be realized by the fine pixels.

  Such features demonstrate unprecedented features in the application of isotope therapy that combines a radiation source and a monoclonal antibody, and allows the dose to be quantitatively grasped while the treatment site is visualized as a high-resolution image. From accurate treatment planning to follow-up. In this way, there is no device that can simultaneously visualize function / morphological information as a diagnostic image, and there is no device that can perform treatment at the same time, which leads to the development of a new isotope preparation. it is conceivable that.

  As a matter of course, this X-ray / gamma-ray imaging apparatus can be used in medical fields, and this detector is also useful in the fields of industrial application and reverse engineering. One example of this is positioning as a non-destructive quantitative analyzer using a detector capable of discriminating and collecting X-rays by energy. The detector can express the linear attenuation coefficients of various energy regions as a CT image, and the elements and compounds constituting the medium can be specified from these values. Also in this regard, a device using this detector can be positioned as a kind of measuring instrument that has never existed until now.

  As described above, the X-ray / gamma ray imaging apparatus according to the present invention eliminates the disadvantages of a single imaging apparatus using X-rays and gamma rays, not only in the medical field but also in industrial use, and is extremely useful equipment. It is.

1 Gantry 2 Bed 12 X-ray tube (radiation source)
13 Detector (Radiation detector)
14 collimator 15 high voltage generator 16 gantry driver 17 detection system driver 18 data collector 21 X-ray controller 22 computer 23 console P subject

Claims (6)

In a radiation detector in which pixels are arranged two-dimensionally, each pixel has a photon counting radiation detection circuit having two or more radiation energy discriminators, and the signal detection information of each pixel and the energy affiliation by the discriminator A cone beam collimator that has an output structure that outputs information independently in units of pixels and directs only to an X-ray beam from an X-ray tube placed opposite to the detector is arranged in front of the detector. An X-ray / gamma ray imaging apparatus, wherein the collimator collimates gamma rays emitted from an object in the field of view and incident on the detector. The X-ray / gamma ray imaging apparatus according to claim 1,
X-ray / gamma-ray imaging characterized in that the X-ray and the gamma-ray are incident at the same time, there is almost no degradation in counting characteristics due to counting down, and the X-ray and the gamma ray are separated and collected by a discriminator. apparatus. The X-ray / gamma ray imaging apparatus according to claim 1,
The X-rays are intermittently irradiated and detected by the detector, and the gamma rays are always emitted from the subject and detected by the detector, and there is almost no deterioration of the counting characteristics due to counting off, and at the time of X-ray imaging, there is almost X An X-ray / gamma ray imaging apparatus characterized in that it can be regarded as a ray transmission image and can be regarded as a gamma ray radiation image when only gamma rays are imaged, and X-rays and gamma rays are separated and collected by a discriminator. The X-ray / gamma ray imaging apparatus according to any one of claims 1 to 3,
An X-ray / gamma-ray imaging apparatus, wherein the detector and the X-ray tube are placed opposite to a subject, and simultaneous tomography of X-ray CT and SPECT is performed by a rotation data acquisition mechanism. The X-ray / gamma ray imaging apparatus according to claim 3,
The position of the X-ray tube and the detector is fixed, and the X-ray image obtained by moving only the subject in the body axis direction, or the position of the detector is fixed without performing X-ray irradiation. An X-ray / gamma ray imaging apparatus characterized in that a gamma ray planar image obtained by moving only the subject in the body axis direction is obtained. The X-ray / gamma ray imaging apparatus according to claim 3,
An X-ray / gamma-ray imaging apparatus characterized in that a gamma-ray planar image obtained without fixing the position of the detector without moving the X-ray and without moving the subject is obtained.

Images