JP2006346290A - Radiographic image photographing apparatus - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a radiographic image photographing apparatus, restraining deterioration of image quality even in a tomographic image or a three-dimensional image, reducing exposure dose, reducing blur of an image due to a motion of a patient due to long photographing time, and decreasing load of a patient due to long time photographing. <P>SOLUTION: This radiographic image photographing apparatus includes: a radiographic image detector 4 for obtaining a phase contrast radiographic image by edge effect when applied radiation is transmitted through an object 3, wherein in the radiographic image detector 4, a radiation detecting element is a one-dimensional or two-dimensional image sensor and generates a tomographic image or a three-dimensional image. The radiographic image detector 4 includes electrodes 8, 11 on both opposite surfaces, and a radiation conversion layer 7 for converting radiation to charges which is provided between the electrodes 8, 11 of both surfaces. The radiation conversion layer 7 is formed of CdTe single crystal or CdZnTe single crystal. <P>COPYRIGHT: (C)2007,JPO&INPIT

Description

  The present invention relates to a radiographic imaging apparatus, and more particularly to a radiographic imaging apparatus that obtains a phase-contrast radiographic image by performing enlarged imaging using a small-focus radiation source to generate a tomographic image or a three-dimensional image. .

  For example, an X-ray CT apparatus has obtained a circular slice image (tomographic image) that cannot be obtained with a conventional simple X-ray photograph, and has seen a breakthrough in image diagnosis. The brain sliced image surrounded by thick bones has greatly improved the diagnostic technology for head injury caused by traffic accidents.

  In recent years, regarding X-ray CT image quality improvement, several studies on X-ray CT using monochromatic X-rays extracted from synchrotron radiation X-rays have been reported. The fundamental important technique here is that a phase-contrast radiation image can be obtained by using monochromatic X-rays, and an attempt has been made to obtain a clear X-ray CT image (Patent Document 1). .

The X-ray CT apparatus is arranged so that X-rays radiated from the radiation tube are narrowed down into a fan-shaped X-ray beam by a collimator at the radiation port, and the radiation tube is opposed to the subject with the subject at the center. The arc-shaped collimator and detector rotate, the detector captures X-ray information transmitted through the subject, and the signal is processed by a computer to obtain a tomographic image of the subject. In recent years, a radiation image detector using a semiconductor single crystal or polycrystal has been proposed as this detector. The radiation image detector uses an X-ray conversion layer made of a semiconductor material that generates charges (electrons-holes) when irradiated with radiation such as X-rays, has high dark resistance, and is resistant to X-ray irradiation. It has a wide dynamic range, good S / N, and good photoconductive properties. For example, Cdte single crystals, CdZnTe single crystals, and the like have been proposed (Patent Document 2).
JP 2002-159486 A JP 2002-202377 A

  By the way, imaging to obtain a phase contrast radiation image by the edge effect when irradiated radiation passes through the subject is an imaging method for obtaining a very high-definition image, but uses a small-focus radiation tube. The radiation output is weaker than that of a normal CT radiation tube, and as long as the current small-focus radiation tube is used, in order to obtain a phase-contrast radiation image, the imaging time tends to be longer in order to obtain the same irradiation dose. is there. In particular, an X-ray CT apparatus or the like originally has a long imaging time, and further imaging for a long time tends to cause blurring of an image due to the movement of the patient, which increases the burden on the patient.

  On the other hand, if a radiograph with a small focal point is used and imaging is performed with a radiographing time similar to the conventional one, the irradiation dose is reduced. Therefore, the use of a conventional detector reduces the image quality of the captured image and hinders diagnosis. The possibility that it will occur will arise. In addition, since the imaging for obtaining the phase contrast radiation image is an enlargement imaging, the amount of radiation reaching the detector is reduced in the first place, and the necessity of a high sensitivity detector is high.

  In the present invention, radiographic imaging capable of reducing exposure dose by suppressing deterioration of image quality even in a tomographic image or a three-dimensional image, reducing blurring of an image due to patient movement due to a long imaging time, and reducing a burden on the patient due to long imaging. A device is provided.

  In order to solve the above problems and achieve the object, the present invention is configured as follows.

  The invention according to claim 1 includes a radiation image detector that obtains a phase-contrast radiation image by an edge effect when irradiated radiation passes through a subject, and the radiation image detector has a one-dimensional radiation detection element or In a radiographic imaging apparatus that is a two-dimensional image sensor and generates a tomographic image or a three-dimensional image, the radiographic image detector has electrodes on both sides facing each other, and converts radiation into charges between the electrodes on both sides. A radiographic imaging apparatus comprising a radiation conversion layer, wherein the radiation conversion layer is made of a CdTe single crystal or a CdZnTe single crystal.

  In the first aspect of the present invention, a good image can be obtained when a tomographic image or a three-dimensional image is generated by obtaining a phase contrast radiation image. In addition, the CdTe single crystal or CdZnTe single crystal of the radiological image detector is a method of imaging by directly converting radiation into electric charges, so there is little reduction in sharpness, and the phase contrast effect that makes the radiation information clear is faithfully imaged, A very clear image can be obtained. Further, since the scattered radiation is removed by the air gap method (takes a longer distance between the subject and the detector) without using a grid in phase contrast imaging, the problem that the apparatus becomes complicated and large in size and difficult to control is eliminated. In addition, radiation does not decrease due to the absorption of direct X-rays (primary X-rays) in the grid, and the SN of the image does not decrease. Even in tomographic images or three-dimensional images obtained by phase contrast imaging, deterioration in image quality is suppressed. It is possible to reduce exposure dose, reduce image blur due to patient movement due to long imaging time, and reduce patient burden due to long imaging time.

  The invention according to claim 2 is the radiographic imaging apparatus according to claim 1, wherein a readout circuit for reading out an output signal from the radiation detection element is a CMOS.

  In the invention according to claim 2, in addition to the high sensitivity by the CdTe single crystal or CdZnTe single connection which is the radiation conversion layer, the use of CMOS capable of pixel image width in the readout circuit further increases the sensitivity. The image quality can be improved and the shooting time can be shortened. In addition, the reading circuit can be miniaturized and an image with high resolution can be output, and a lighter and more compact configuration can be achieved.

  In the invention according to claim 3, the focal diameter of the radiation tube for irradiating the radiation is D, the distance from the radiation tube to the center of the subject is R1, and the distance from the center of the subject to the radiation image detector is 3. The radiographic imaging apparatus according to claim 1, wherein when the distance is R2, 0.25 m ≦ R1 ≦ 1.5 m, 0.25 m ≦ R2 ≦ 1.5 m, and 10 μm ≦ D ≦ 300 μm. is there.

  In the invention according to claim 3, the focal diameter of the radiation tube is defined by a distance from the radiation tube to the center of the subject, and a distance from the center of the subject to the radiation image detector. A good image can be obtained when generating a tomographic image or a three-dimensional image.

  Hereinafter, embodiments of the radiation image capturing apparatus of the present invention will be described in detail with reference to the drawings. However, the present invention is not limited to the embodiments.

  An embodiment in which the present invention is applied to an X-ray CT apparatus capable of obtaining a phase-contrast radiation image will be described in detail.

[Phase contrast radiation image]
First, phase contrast will be described. As schematically shown in FIG. 1, X-rays as radiation are a kind of electromagnetic waves and have wave properties. Hereinafter, radiation is also referred to as X-ray, and X-ray is a preferred embodiment.

  Similar to visible light, when X-rays pass through the subject 3 which is an object having a different refractive index, refraction occurs at the interface. In the X-ray transmission image on the radiation image detector 4 at the interface portion having a different refractive index, the X-ray intensity is reduced by the refraction of X-rays, and the refracted X-rays overlap with the X-rays traveling straight through the space. Thus, there is a portion where the X-ray intensity increases. That is, in the negative image obtained here, a portion where the X-ray intensity decreases at an interface having different refractive indexes is whitened and a portion where the X-ray intensity increases is blackened, so that a so-called edge-enhanced image is obtained. This is a phenomenon called phase contrast. Since the wavelength of X-ray is very short and its refractive index is small, this phase contrast has been overlooked in conventional X-ray imaging.

  In conventional X-ray images, images based on phase contrast (phase contrast radiation images) are not fully utilized, but rather X-ray images (absorption contrast radiation images) with only absorption contrast due to X-ray absorption differences have been used. .

  In the present invention, an X-ray CT apparatus for photographing a high-quality X-ray CT image with good sharpness by this phase contrast is provided. Here, a method using a synchrotron X-ray beam such as the High Energy Laboratory in Tsukuba City or SPring-8 in Hyogo Prefecture is well known for the phase contrast radiation image. It is obvious that synchrotrons are too large to be used widely in general medical facilities. Further, according to the method described in JP-A-11-502620, the X-ray focal spot size is very small, and the X-ray source and the film are separated from each other by a distance that can be regarded as a point focal point. It has been reported to take images. However, with this method, the X-ray becomes too weak and it is impossible to take an X-ray image of the human body.

  In order to obtain an X-ray phase contrast radiation image, in principle, the distance R2 from the subject 3 to the radiation image detector 4 must be a certain distance as shown in FIG. In addition, the radiation tube 1 used for X-ray CT has to take a value of 50 μm or more instead of a focal point of the radiation tube in order to obtain a constant X-ray intensity. Here, the focal diameter is the size of the window where X-rays are emitted from the radiation tube, and generally has a square shape and the length of one side thereof. In such a case, a blur which is geometrically sharp due to the focal diameter is generated. The size B of blur is expressed by B = a × (R2 / R1). Since synchrotron X-rays obtained from the synchrotron are close to parallel lines, that is, equivalent to a focal point of infinity (R1 → ∞), the blur size B is 0, and the problem of such blur is Absent. Further, when using a microfocus X-ray source, a → 0, and automatically B → 0. In this case, it is not necessary to consider the influence of blur.

  Here, in X-ray CT imaging, as shown in FIG. 3, the radiation tube 1 and the radiation image detector 4 are arranged around the subject (subject 3), and the radiation tube 1 and the radiation image detector 4 are synchronized. Thus, an image is taken by rotating around the subject (subject 3). In order to obtain a phase contrast radiographic image effectively, the focal diameter and the rotation radius length of the radiation tube must be in a constant region.

  In the X-ray CT, the radiation tube 1 and the radiation image detector 4 are simultaneously rotated around the subject (subject 3) as described above. In this case, the turning radius is naturally limited. That is, the smallest radius that does not hit the subject (subject 3) is about 0.5 m. On the contrary, when the rotation radius is increased, the X-ray CT apparatus itself is increased, and the largest radius is about 2 m.

  In addition, the focal diameter of a radiation tube said here can be measured using a pinhole camera or a test chart as defined in JISZ4702. In this embodiment, cone beam X-rays are emitted from a radiation tube, and this is detected by a planar radiation image detector, so that the imaging time is short and the X-ray CT image has high image quality due to phase contrast. Can be obtained.

  In order to obtain a phase contrast radiographic image, the focal diameter of the radiation tube that irradiates the radiation is D, the distance from the radiation tube to the center of the subject is R1, and the distance from the center of the subject to the radiation image detector is R2 is preferably 0.25 m ≦ R1 ≦ 1.5 m, 0.25 m ≦ R2 ≦ 1.5 m, and 10 μm ≦ D ≦ 300 μm. When a phase contrast radiation image is obtained and a tomographic image or a three-dimensional image is generated. A good image is obtained.

Furthermore, 0.5m ≦ R1 ≦ 1m is preferable. The longer R1, the greater the phase contrast effect and the clearer the image (the smaller the geometrical unsharpness of the focal spot diameter as in the synchrotron described above). On the other hand, if R1 is too long, the apparatus becomes larger and the usability becomes worse. Further, 0.5 m ≦ R 2 ≦ 1 m is preferable. Since X-ray refraction occurs at the subject position and gradually increases toward the radiation image detector, the longer R2, the greater the phase contrast effect, and the clearer the image. On the other hand, if R2 is too long, the apparatus becomes large and the usability becomes poor, and the geometrical insensitivity due to the focal diameter D becomes large and the image is blurred. Further, it is preferable that 50 μm ≦ D ≦ 200 μm, and more preferably 70 μm ≦ D ≦ 120 μm. Although the X-ray output increases as the focal diameter D is larger (although the irradiation X-ray dose per unit time increases), if it is too large, the image due to geometrical sharpness will be blurred. On the other hand, the smaller the focal diameter D, the smaller the geometrical insensitivity and the clearer the image becomes. However, when the focal diameter D is too small, the X-ray output is small and practical X-ray imaging cannot be performed. Each of 0.5 m ≦ R 1 ≦ 1 m, 0.5 m ≦ R 2 ≦ 1 m, 50 μm ≦ D ≦ 200 μm, and 70 μm ≦ D ≦ 120 μm can be arbitrarily selected and set according to the apparatus and photographing.
[Radiation image detector]
(Radiation detection element)
The radiation image detector 4 is a one-dimensional or two-dimensional image sensor with a radiation detection element, and this embodiment is shown in FIG. FIG. 4 is a principle diagram of a direct conversion X-ray image sensor. The basic configuration of the image sensor is a radiation conversion layer 7 that converts X-rays into electric charges, a storage capacitor (capacitor) 12 that stores electric signals from the radiation conversion layer 7 as electric charges, and charge storage / voltage output of the storage capacitors 12. A readout circuit 13 for reading out signals, a bias electrode 8 for applying a bias voltage V to the radiation conversion layer 7, and a pixel electrode 11 are included. A bias voltage V is applied so that the bias electrode 8 has a positive potential and the pixel electrode 11 has a negative potential. The sign of the applied potential varies depending on the material of the radiation conversion layer 7.

  The pixel electrode 11, the radiation conversion layer 7, and the bias electrode 8 are laminated, and the storage capacitor 12 and the readout circuit 13 are laminated so as to cover almost the entire surface on the pixel electrode 11 side of this one-dimensional or two-dimensional array. ing. Among the X-rays, X-rays that are transmitted without being absorbed by a subject such as a human body enter the radiation conversion layer 7. In the direct conversion method, charges (hole-electron pairs) corresponding to the X-ray dose incident on the radiation conversion layer 7 are excited.

  A bias voltage V is applied to the radiation conversion layer 7 by a bias electrode 8 formed on the entire surface and a pixel electrode 11 formed for each cell. The generated carriers move according to the polarity of the bias applied to the radiation conversion layer 7, electrons move to the bias electrode 8, holes move to the pixel electrode 11, and are stored as charges in the storage capacitor (Cs) 12 in the array. . The electrode of the radiation image detector 4 is an electrode on the signal extraction side, that is, the pixel electrode 11 is separated for each pixel, and the electrode on the opposite side (usually the direction in which X-rays are incident), that is, the bias electrode 8 is a common electrode (one sheet). Electrode).

  In the present invention, the radiation conversion layer 7 is made of a CdTe single crystal or CdZnTe single crystal, and this structure will be described below. As shown in FIG. 5, in the radiation detection element, the radiation conversion layer 7 is composed of a CdTe or CdZnTe semiconductor substrate, and a plurality of bias electrodes 8 and pixel electrodes 11 are formed on these two surfaces. The portions other than the portions where the electrodes are formed are covered with an insulator film 14 made of aluminum, nitrogen, and oxygen, and the surface of the semiconductor substrate does not come out directly.

  As shown in FIG. 6, it is more preferable that the insulator film 14 has a structure in which a hole is formed only on a portion where the bump material is in contact with the pixel electrode 11 as shown in FIG. It is possible to prevent the bump material from penetrating the surface of the semiconductor substrate from between the pixel electrode 11 and the insulator film 14. On the contrary, as shown in FIG. 7, it is preferable that the pixel electrode 11 is in contact with the surface of the semiconductor substrate at the center, and the peripheral portion overlaps the insulator film 14. The shape of the pixel electrode 11 includes a circle, a rectangle, a square, a rectangle with rounded corners, and the like. In such a case, the pixel electrodes 11 are arranged in rows and columns as shown in FIG. It is normal.

  Below, an example of the manufacturing process of a CdTe radiation detection element is demonstrated.

  As shown in FIG. 9, first, In was formed as a bias electrode 8 on the back surface of the CdTe substrate by vacuum deposition (FIG. 9A). Next, a photoresist 20 is applied to the opposite surface (FIG. 9B), the photoresist is exposed through a photomask on which a pixel electrode pattern is drawn, and further developed to form a pixel of the photoresist. An opening 20a was opened in the electrode forming portion (FIG. 9 (c)). This was put in a chloroplatinic acid aqueous solution, and after forming a platinum electrode 21 by an electroless plating method, a gold thin film 22 was further deposited by a vapor deposition apparatus (FIG. 9 (d)). Thereafter, the photoresist was washed by washing with an acetone solution to form a pixel electrode 11 on the surface of the CdTe substrate (FIG. 9E).

  Next, an insulating film 14 made of an amorphous aluminum oxynitride thin film was formed using an inductively coupled high-frequency plasma assisted magnetron sputtering apparatus (FIG. 9 (f)). Aluminum was used as a target, and a 10% oxygen mixed nitrogen gas was introduced into the sputtering apparatus as a reaction gas. Argon was also introduced for plasma generation. The film forming pressure was 0.1 Pa. This was taken out after forming into a film thickness of 100 nm. Photoresist 20 was applied again thereon (FIG. 9G), developed after exposure with a mask having an electrode opening pattern, and a hole of electrode opening 20b was formed in the photoresist (FIG. 9H )). Further, the amorphous aluminum oxynitride thin film was etched with an aqueous solution of sodium hydroxide to expose a part of the upper part of the electrode (FIG. 9 (i)). Thereafter, the photoresist was dissolved and washed with acetone to obtain a radiation detection element (FIG. 9 (j)).

  This was flip-chip bonded to the readout integrated circuit constituting the storage capacitor 12 and readout circuit 13 to constitute the radiation image detector 4. When an object was placed between the radiation image detector 4 and the X-ray source, X-rays were irradiated and the image was copied to a computer, a good image free from defects was obtained.

  10 and 11 are a side view and a front view showing the unit image detector 31 constituting the radiation image detector 4. FIG. In the unit image detector 31, the radiation detection element 32 is connected to the readout circuit 13 in a pixel unit by a flip chip bonding method. The semiconductor radiation detection element 32 and the readout circuit 13 are joined to each other as shown in the figure to form a rectangular unit image detector 31 having a length of 115 mm and a width of 45 mm, for example. The readout circuit 13 has a signal extraction pad 36 on one side, and is directly bonded to the radiation detection element 32 by a flip chip bonding method so that the signal extraction pad 36 portion protrudes. When used as a one-dimensional detector, the size of the unit image detector may be about 25 mm × 10 mm.

  Then, the two unit image detectors 31 collide with each other on the side opposite to the signal drawing pad 36 side without any gap, and the unit image detectors 31 that are confronted with each other as a set, A radiographic image detector is configured by sequentially colliding a plurality of sets with a side surface on which the signal extraction pad 36 is not provided and arranging them in parallel on a rectangular substrate 37 as shown in FIGS. .

  FIGS. 12 and 13 show a case in which 10 unit image detectors 31 in total, 2 in the vertical direction and 5 in the horizontal direction, are arranged on the substrate 37 and wire-bonded to form a radiation image detector 4 having an effective field of view of 230 mm × 225 mm. Indicates. The signals from the unit image detectors 31 are sequentially supplied to an external image configuration unit and an image display unit via the substrate 37. In addition, the radiation image detector 4 of arbitrary magnitude | size can be comprised by increasing / decreasing the number of collisions of the unit image detector 31 and the assembly | attachment number of the radiation image detector 4 which consists of several unit image detectors 31. .

  The radiation image detector 4 may be an area sensor or a line sensor. The area sensor can acquire more image information within the same time, and if the amount of information that can be obtained is the same, it is possible to perform imaging in a short time (small irradiation dose). On the other hand, the line sensor can make the apparatus compact and can produce a detector that is long in the one-dimensional direction, so that the entire imaging area can be increased.

The pixel size of the radiation image detector 4 is preferably 50 μm or more and 200 μm or less, more preferably 80 μm or more and 160 μm or less. When the pixel size is reduced, the fill factor is reduced and the SN of the image is deteriorated. On the other hand, if the pixel size is too large, the amount of information is small and the image is smoothed, which causes a problem as a diagnostic image. The size of the radiation image detector 4, that is, the imaging size is preferably 10 × 10 cm or more, and more preferably 20 cm × 20 cm or more. If it is not large enough, it will interfere with the diagnosis. From the current use, it may be about 45 cm × 45 cm, and about 30 cm × 30 cm is sufficient for diagnosis.
(Read circuit)
Next, a preferred embodiment of the readout integrated circuit will be described in detail. FIG. 14 is a configuration diagram in which the readout circuit 13 of the readout integrated circuit is composed of CMOS. The readout circuit 13 includes an electrode 110 having a CMOS 130 formed in a matrix (a CMOS array 131 composed of n1 columns of CMOS 1330 is arranged in n2 rows) and an A / D converter 121 corresponding to each CMOS array 131. Includes a signal processing unit 120 arranged in n2 rows.
The electrode 110 is bonded to each pixel of each element of the radiation detecting element 32 in a one-to-one bonding manner, and is read as a signal of each pixel.

  Each A / D converter 121 includes an amplifying unit 13 and an A / D converting unit 14, and amplifies the output of the CMOS 13O and converts it into digital data of 12 to 16 bits (4096 to 65536 gradations). The read charge information is converted into a digital image signal and sequentially output.

In the present invention, the readout circuit 13 for reading out an output signal is preferably a CMOS (complementary metal-oxide semiconductor), and pixel amplification is possible with this CMOS (E.Oba, K. Mabuchi). , Y.Iida.N.Nakamura, and H.Mimura: A 1/4 Inch 330 K Square Pixel Progressive Scan CMOS Active Pixel Image Sensor, ISSCC Dig. Of Tech. Papers, pp. 180-181 (1997) JED Fuwitz, PBDenyer, DJ Baxter and G. Townsend: An 800k-Pixel Color CMOS For Consumer Still Camera, SPIE Vol. 3019, pp. 115-124 (1997)). In addition, the shooting time can be shortened. Further, in addition to the CdTe single crystal or CdZnTe single crystal of the radiation image detector 4, a further high sensitivity detector can be obtained by CMOS amplification. In addition, the reading circuit can be miniaturized and a high-resolution image can be output, and a lighter and more compact configuration can be achieved.
[Photographing device configuration]
This radiographic imaging device may be a CT dedicated device or a C-arm imaging device. FIG. 15 is a diagram illustrating a configuration example of an imaging apparatus using a C-arm.

  The radiographic imaging device includes a detector vertical motion drive unit 301, a detector longitudinal motion drive unit 302, a radiation image detector 4, a C arm longitudinal motion vertical motion drive unit 311, a C arm rotation drive unit 339, an arm 315, C An arm slide rotation drive unit 316, a main body 317, an X-ray diaphragm unit 324, a C arm 325, the radiation tube 1, and the like are included.

  The detector vertical movement drive unit 301 moves the detector front / rear left / right movement drive unit 302 and the entire detector 306 up and down. The radiation image detector 4 can be moved up and down in a compact manner by using a plurality of cylinders with different thicknesses provided with a mechanism that can be expanded and contracted in the detector vertical movement drive unit 301.

The radiation image detector 4 is held by the detector back-and-forth left-right motion drive unit 302, and the radiation image detector 4 can be freely moved in a plane orthogonal to the center of the X-ray bundle. The detector vertical movement mechanism 301 and the radiation tube 1 are fixed to both ends of the C arm 325. The slide angle of the C arm 325 allows the X-ray irradiation angle to be freely adjusted while keeping the center of rotation coincident with the region of interest diagnosis. This C-arm imaging apparatus is preferable because it can perform simple X-ray imaging in addition to tomography, and it is not necessary to move between the CT apparatus and the general X-ray imaging apparatus.
(Generation of 3D images)
In the above embodiment, the radiographic imaging apparatus that generates a tomographic image by X-ray CT imaging has been described. Hereinafter, the radiographic imaging apparatus that generates a three-dimensional image will be described. In X-ray CT imaging, the radiation tube 1 and the radiation image detector 4 may be a CT system that rotates once around the subject (360 degrees), or reconstructs an image as tomosynthesis with a rotation of 360 degrees or less. Also good. The CT method is a large amount of information and is considered useful for diagnosis, but the tomosynthesis method is convenient because the device is compact and the imaging time can be shortened.

  A tomosynthesis system to which this tomosynthesis system is applied is shown in FIG. FIG. 16 schematically shows a front image of the tomosynthesis system 201 used to form a three-dimensional image of the subject 202. FIG. 17 shows a side view of the system 201, and FIG. 18 shows a top view of the system 201.

  The system 201 includes an X-ray source 211. The X-ray source 211 preferably includes a radiation tube disposed inside the housing. The X-ray source 211 shown in FIG. 16 can rotate along an arcuate path in one plane. In addition to the X-ray source 211 rotating about the rotation axis 212 along the arcuate path 204, the X-ray source 211 moves in the radial direction 205 and moves out of the plane of rotation in the direction 206. You can also The rotary shaft 212 can have various heights above the radiation detector 4, but is preferably close to the height of the surface of the radiation detector 4. The X-ray source irradiates the subject 2020 with an X-ray beam 203 from a plurality of positions in the sector (that is, from some, but not all, positions along the path in the sector).

The radiographic image detector 4 is positioned with reference to the X-ray source 211 so as to detect the X-ray 203 transmitted through the subject 202. The system 201 further creates an image to be imaged based on the X-ray 203 detected by the radiation image detector 4 from the scan of the X-ray source 211 in the sector.
[Image output format]
The obtained image is displayed on a CRT or a liquid crystal image display device. Furthermore, it is drawn as a hard copy image by a silver salt film, a sublimation type, an ink jet printer, or the like, and is observed on Syaukasten. This image may be a monochrome image display or a color image using pseudo color.

  The obtained image signal can be stored for the purpose of temporary or permanent storage. At this time, a magneto-optical disk or a DVD can be used. In addition, it is a preferable aspect that an image signal is transmitted to a remote place using a line and an image is displayed at the place to perform image diagnosis.

  In addition, since this invention can obtain a sharp image using phase contrast, it is suitable also when the fine blood vessel in a brain is made into imaging | photography object. In this case, since the amount of contrast agent is smaller than that in the conventional case, a secondary effect is obtained in that the safety of the subject at the time of tomography is increased.

  In addition, a tomographic image, a three-dimensional image, and a transmission image can be regenerated from the image data obtained from the apparatus of the present invention. The regenerated image can be output to a monitor or a printer, but it is preferable to output it on a monitor from the viewpoint that three-dimensional display and a large number of images can be switched efficiently. It can also be used with still images.

  Thus, in the present invention, a good image can be obtained when a phase contrast radiation image is obtained and a tomographic image or a three-dimensional image is generated. In addition, the CdTe single crystal or CdZnTe single crystal of the radiation image detector 4 is a method of imaging by directly converting radiation into electric charges, so there is little reduction in sharpness, and the phase contrast effect that makes radiation information clearer is faithfully imaged. A very clear image can be obtained. Further, since the scattered radiation is removed by the air gap method (takes a longer distance between the subject and the detector) without using a grid in phase contrast imaging, the problem that the apparatus becomes complicated and large in size and difficult to control is eliminated. In addition, radiation does not decrease due to the absorption of direct X-rays (primary X-rays) in the grid, and the SN of the image does not decrease. Even in tomographic images or three-dimensional images obtained by phase contrast imaging, deterioration in image quality is suppressed. It is possible to reduce exposure dose, reduce image blur due to patient movement due to long imaging time, and reduce patient burden due to long imaging time.

  The present invention can be applied to a radiographic imaging apparatus that generates a tomographic image or a three-dimensional image by obtaining a phase-contrast radiation image by performing magnified imaging using a small focal radiation source. It is possible to reduce the exposure dose by suppressing the decrease, to reduce the blur of the image due to the movement of the patient due to the long imaging time, and to reduce the burden on the patient due to the long imaging time.

It is a block diagram which shows the structure of a phase contrast X-ray CT apparatus. It is explanatory drawing for the principle explanation of a phase contrast radiographic imaging device. It is explanatory drawing for the principle explanation of a phase contrast radiographic imaging device. It is a principle block diagram of a direct conversion type X-ray image sensor. 1 is an overall perspective view of a semiconductor radiation detection element according to an embodiment. It is principal part sectional drawing which shows the state in which the insulator film overlapped on the electrode. It is principal part sectional drawing which shows the state in which the peripheral part of the electrode has overlapped on the insulator film. It is a top view which shows the example of an arrangement | sequence of a pixel electrode. It is a figure which shows the manufacture process of the CdTe radiation detection element for X-ray imaging devices. It is a side view which shows the unit image detector of the radiographic image detection apparatus by one Embodiment. It is a front view of the unit image detector shown in FIG. It is a front view which shows a radiographic image detector. FIG. 13 is a side view of the radiation image detector shown in FIG. It is a circuit diagram of a read circuit. It is a side view of a mobile X-ray imaging apparatus. It is a schematic diagram of the front image of the system according to the embodiment. It is a schematic diagram of the side view of a system. It is a schematic diagram of the upper surface image of a system.

Explanation of symbols

3 Subject
4 Radiation image detector
7 Radiation conversion layer
8 Bias electrode
11 pixel electrodes

Claims (3)

A radiation image detector that obtains a phase-contrast radiation image due to an edge effect when irradiated radiation passes through the subject;
The radiological image detector is a radiographic imaging device in which a radiation detection element is a one-dimensional or two-dimensional image sensor and generates a tomographic image or a three-dimensional image.
The radiation image detector has electrodes on both sides facing each other,
A radiation conversion layer for converting radiation into electric charge between the electrodes on both sides;
The radiation image capturing apparatus, wherein the radiation conversion layer is made of a CdTe single crystal or a CdZnTe single crystal.   2. The radiographic image capturing apparatus according to claim 1, wherein a readout circuit that reads an output signal from the radiation detection element is a CMOS. The focal diameter of the radiation tube for irradiating the radiation is D,
The distance from the radiation tube to the center of the subject is R1,
When the distance from the center of the subject to the radiation image detector is R2,
3. The radiographic imaging apparatus according to claim 1, wherein 0.25 m ≦ R 1 ≦ 1.5 m, 0.25 m ≦ R 2 ≦ 1.5 m, and 10 μm ≦ D ≦ 300 μm.