JP2006218327A - Radiation imaging apparatus - Google Patents

Radiation imaging apparatus Download PDF

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JP2006218327A
JP2006218327A JP2006144354A JP2006144354A JP2006218327A JP 2006218327 A JP2006218327 A JP 2006218327A JP 2006144354 A JP2006144354 A JP 2006144354A JP 2006144354 A JP2006144354 A JP 2006144354A JP 2006218327 A JP2006218327 A JP 2006218327A
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radiation
imaging
dimensional detector
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radiation source
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Osamu Tsujii
修 辻井
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Canon Inc
キヤノン株式会社
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<P>PROBLEM TO BE SOLVED: To implement various measures against the problem that it is difficult to properly determine a distance between a two-dimensional detector and a radiation generating source in a rotary CT unattended with a spiral. <P>SOLUTION: This radiation CT imaging apparatus comprises the radiation generating source (101) for irradiating a subject with radiation, a rotating means (102) for rotating a human body in the radiation, and the two-dimensional detector (105) for detecting the radiation, and a distance from the focus of the radiation to the two-dimensional detector (105) is approximately 200 cm or more. <P>COPYRIGHT: (C)2006,JPO&NCIPI

Description

  The present invention relates to a radiation imaging apparatus that captures a three-dimensional image using a cone beam as a radiation generation source, and more particularly to a radiation imaging apparatus that suitably selects a distance between the radiation generation source and a two-dimensional detector.

  In recent years, in order to acquire digital data on a large screen, development of a two-dimensional detector for radiography (also referred to as FPD (Flat Panel Detector)) is in progress (for example, see Patent Document 1). In particular, an imaging device using a two-dimensional detector having a large light receiving surface of 43 cm × 43 cm is in practical use for simple imaging.

  On the other hand, X-rays are exposed to the subject, X-rays transmitted through the subject are detected by an X-ray detector, and the subject is seen through based on this X-ray detection output (number of photons of X-rays). An X-ray CT apparatus that captures an image (referred to as a scanogram or SCOUT image), a tomographic image, or a three-dimensional image is known.

  In this X-ray CT apparatus that captures a three-dimensional image, the development technology of a two-dimensional detector has been improved, and a cone beam CT apparatus has been developed as an X-ray CT apparatus that captures a three-dimensional image. In a normal X-ray CT apparatus, an X-ray beam is cut out thinly in the Z direction and is called a fan beam. In a cone beam CT (hereinafter referred to as “CBCT”), an X-ray beam (also called “CBCT”) that spreads in the Z direction ( Hereinafter, this cone beam is received by a two-dimensional detector. Compared with CT using a fan beam, cone beam CT has a wider range in which a subject can be imaged by scanning with one rotation, and therefore has the advantage of reducing the number of rotations and improving the efficiency of imaging. In other words, it is possible to improve the efficiency of photographing by taking a wide cone angle, which is the spread of the cone beam. However, on the other hand, if the cone angle is too wide, there is a problem that a reconstruction error occurs in the reconstructed image.

  By the way, cone beam CT includes a type in which a pair of an X-ray source and a detector scans (collects projection data) while rotating around the subject (for example, Patent Document 2). However, in order to irradiate the radiation from the radiation generating source to all the light receiving surfaces of the two-dimensional detector having a wide light receiving surface with a certain cone angle or less, the distance between the two-dimensional detector and the radiation generating source must be set to a certain value or more. There is a problem that must be. Therefore, in a type of CBCT in which scanning (collection of projection data) is performed while a pair of an X-ray source and a detector rotates around the subject, the distance between the two-dimensional detector and the radiation generation source is effective for a large light receiving surface. Arranging for use is difficult due to the size of the device.

  On the other hand, an object rotation type CBCT in which a pair of an X-ray source and a detector is fixed and the object rotates instead (no spiral) is being developed for practical use (for example, Patent Document 3). .

  In simple X-ray imaging, there is a background in which the distance between the subject and the radiation source has been determined by trial and error in the tradition of about 100 years, and the distance between the subject and the radiation source should be determined appropriately. Is one of the good and bad skills of X-ray technicians. As described above, it is important for the radiography apparatus to appropriately determine the distance between the two-dimensional detector and the radiation generation source, and the subject rotation type CBCT in which the subject rotates, in which the X-ray technician has no imaging experience, is put into practical use. In order to achieve this, it is considered particularly important to appropriately determine the distance between the two-dimensional detector and the radiation source.

  For example, in CBCT disclosed in Patent Document 3, X-ray I.D. I. Is set to 1200 (mm), the distance from the focal position of the X-ray generator to the rotation center of the rotating device, that is, the rotation radius r of the X-ray source is set to 800 (mm). . I. Is 16 inches (the horizontal screen size is 400 (mm)), the field of view of the transmitted X-ray image is a spherical shape having a diameter of about 260 (mm). When calculated from this condition, the cone angle is 9.5 degrees on one side and 19 degrees in total. However, there is a problem that a reconstruction error occurs in the reconstructed image in the peripheral area of the visual field (mainly the peripheral area of the two-dimensional detector). There is a problem that an appropriate cone angle is not required.

  In addition, the spiral type (for example, Patent Document 2) sets the cone angle of CBCT to be relatively small (1-2 degrees), and collects data of the entire target region by a plurality of rotations. Examination of the cone angle at which a reconstruction error occurs is not made in the spiral CBCT.

However, in the subject rotation type CBCT in which the pair of the X-ray source and the detector is fixed and the subject rotates instead, there is a problem that the distance between the two-dimensional detector and the radiation generation source is not appropriately determined. In CBCT, there is a problem that a reconstruction error of a reconstructed image occurs, and there is a problem that an appropriate cone angle is not obtained.
JP 09-288184 A Japanese Patent Laid-Open No. 10-21372 JP 2000-217810 A

  Conventionally, since it has been difficult to appropriately determine the distance between the two-dimensional detector and the radiation source, various countermeasures have been desired.

  Therefore, an object of the present invention is to obtain a reconstructed image in which the distance between the two-dimensional detector and the radiation source is appropriately determined in CBCT and the reconstruction error is small.

  A radiation source for exposing a subject to radiation, a rotating means for relatively rotating a subject in the radiation emitted by the radiation source, a two-dimensional detector for detecting the radiation, and the two-dimensional detector A radiographic apparatus having a reconstructing means for reconstructing an output signal from the image forming apparatus, having two photographing modes, wherein the first mode is a plurality of images by the two-dimensional detector while rotating the subject by the rotating means. To acquire a three-dimensional image of the subject by reconstructing the plurality of images, and in the second mode, the image is captured by the two-dimensional detector without rotating the subject, and the image is not reconstructed. A two-dimensional image of a subject is acquired. Further objects and other features of the present invention will become apparent from the preferred embodiments described below with reference to the accompanying drawings.

  As described above, according to the present invention, the distance between the two-dimensional detector and the radiation source can be appropriately determined in CBCT, and a reconstructed image with few reconstruction errors can be obtained.

  FIG. 1 shows a configuration example of a subject rotation type CBCT in which a subject rotates, where FIG. 1A is a top view and FIG. 1B is a side view. In the figure, reference numeral 101 denotes a radiation source that emits cone beam radiation toward a subject, and at the same time, a focal point of the radiation. Here, as shown in the figure, the spread angle in the vertical direction of radiation is called a cone angle, and the spread in the horizontal direction of radiation is called a fan angle. Reference numeral 102 denotes a rotary table that rotates with a subject 106 irradiated with radiation emitted from the radiation source 101. Reference numeral 104 denotes a breastplate for fixing a subject, which is supported by a column 103 fixed to a rotary table. Reference numeral 105 denotes a two-dimensional detector. Radiation emitted from the radiation source 101 passes through a breastplate 104, a subject 106, and a scattered radiation removal grid (not shown) that fix the subject, and is detected by the two-dimensional detector 105 to be an electrical signal. To be changed.

  The two-dimensional detector 105 is composed of, for example, a semiconductor sensor. For example, one pixel is composed of 250 × 250 μm, and the sensor outer shape is composed of 43 × 43 cm. In this case, the number of pixels is 1720 × 1720 pixels.

  Further, in FIG. 1, the distance between the focal point 101 and the rotation center of the rotary table 103 (Focus-Center-Distance) is abbreviated as FCD, and is hereinafter referred to as FCD. (Focus-Detector-Distance) is hereinafter abbreviated as FDD. In addition, a real space range in which a three-dimensional image is reconstructed by CBCT is referred to as a reconstruction region, and is usually a cylindrical region. The height of this reconstruction area is referred to as “reconstruction height (Height of View)” and abbreviated as “HOV” hereinafter. The radius of the reconstruction area is called an effective field diameter (Field of View), and is hereinafter abbreviated as FOV.

  FIG. 2 is a diagram showing an example of a system configuration diagram of the present invention. In FIG. 2, 201 is a BUS, and control signals and data are transmitted and received via this BUS. 202 is a CPU if it is a computer, and is a control means for controlling the entire system. 203 is an information input means for inputting subject information of a subject's height, weight, and imaging region (chest, abdomen, head, etc.). In particular, in the case of chest imaging, subject information peculiar to each imaging region is also input from the information input means, such as inputting the height of lung field height (lung field height). Reference numeral 204 denotes interface means for giving an instruction regarding the operation of the system such as an instruction to start photographing or an emergency stop of the system. Reference numeral 205 denotes calculation means for calculating a suitable FDD distance between the radiation source 101 and the two-dimensional detector 105 based on the subject information input by the information input means 203, and 206 denotes radiation at the distance calculated by the calculation means 205. It is a radiation source moving means for arranging the generation source 101 and the two-dimensional detector 105 using a moving mechanism including a motor (not shown). Either the radiation source 101 or the two-dimensional detector 105 may be moved in the arrangement, but generally the radiation source 101 is moved, and the radiation source moving means 206 is driven by the control of the control means 202. Is done.

  Reference numeral 207 denotes a rotary table motor control unit that rotates the rotary table 102 under the control of the control unit 202 in accordance with a shooting start instruction input by the interface unit. Reference numeral 208 denotes a radiation source output control unit that controls the radiation state of the radiation source 101, such as start and end of radiation, under the control of the control unit 202 in accordance with an imaging start instruction input by the interface unit. . Reference numeral 210 denotes reconstruction means for constructing a reconstructed image (a cross-sectional image can be obtained from the reconstructed image) from the output image of the two-dimensional detector 105. A display unit 211 displays the image reconstructed by the reconstruction unit 210 or displays the distance calculated by the calculation unit 205.

  3 and 4 are flowcharts showing the processing flow of the radiation imaging apparatus. The operation of the subject rotation type CBCT in which the subject rotates according to this processing flow will be described by taking chest imaging as an example.

  First, according to the flow of FIG. 3, an FDD that is a distance (Focus-Detector-Distance) between the focal point (which is also the position of the radiation generation source) 101 and the two-dimensional detector 105 is preferably calculated, and the focal point (radiation generation source) is calculated. The flow of processing for arranging the 101 and the two-dimensional detector 105 on the calculated FDD will be described.

  An operator such as an X-ray engineer inputs patient information such as the height and weight of the subject via the information input unit 202 before starting imaging (S301). Next, the calculation means 205 statistically calculates the lung field height (the length of the lung field in the height direction) from the height and weight included in the patient information. The calculation means 205 has a conversion table for converting the height of the lungs from the height and weight, and the height of the lungs is calculated from the height and weight. In general, height and lung field height are highly correlated, and it is possible to estimate the lung field height from existing statistics. Furthermore, by adding the patient's weight as information, the lung field height can be estimated more accurately. It is because it is possible.

  In chest radiography, the lung field is mainly used as a diagnostic area for the doctor, so that the lung field height is radiographing that matches the HOV that is the height of the reconstruction area. Therefore, the estimated value of the lung field height is determined as the HOV value (S302). In particular, when it is desired to photograph accurately, an engineer can measure and input the lung field height from the appearance of the patient before photographing. In this case, the measurement value becomes HOV as it is.

  This is because, particularly in the case of chest imaging, in the chest imaging, the lung field is mainly the target region for diagnosis by the doctor, and therefore it is desirable that the lung field height be an imaging that matches the above-mentioned HOV.

Next, the calculation means 205 calculates FDD which is the distance (Focus-Detector-Distance) between the focal point (which is also the position of the radiation source) 101 and the two-dimensional detector 105 using the equation (1) from the determined HOV. Calculate (S303). Here, the calculation means 205 determines the cone angle φ between approximately 6 ° and 10 ° from the imaging part information obtained by the information input means. For a part that has a fine structure such as the chest and is also an important diagnostic target region around the lung field, a high-resolution image is required for the entire reconstruction region, so the cone angle φ is approximately 10 ° or less. select. On the other hand, when the reproducibility of a fine structure is not strictly required as in the abdomen, 10 ° or more may be selected. In this case, since the FDD is shorter than when 10 ° is selected, there is an effect that the radiation dose of the radiation source 101 can be suppressed. As the cone angle is reduced, a high-resolution reconstructed image can be obtained, but the FDD becomes longer and a radiation source having a large anode heat capacity is required. On the other hand, as the cone angle is increased, the FDD is shortened, and the anode heat capacity of the radiation generating source can be reduced. However, the size of the cone angle is naturally limited by the required resolution of the reconstructed image as described above.
FDD = 0.5 * FOV + 0.5 * HOV / tan (φ / 2) [mm] (1)
This equation (1) is calculated from experiments and clinical trials, and is used to favorably determine the FDD. A focal point (even at the position of the radiation source) is set at a distance of FDD or more calculated with a cone angle φ of 10 ° or less. It was made clear that an appropriate reconstructed image resolution can be obtained if the distance between 101 and the two-dimensional detector 105 is maintained. Detailed description will be given later. At the same time, the FDD calculated by the image display unit 211 is displayed (S303).

  In general, it is necessary for an X-ray technician to determine the FDD which is the distance (Focus-Detector-Distance) between the focal point (which is also the position of the radiation source) 101 and the two-dimensional detector 105. Since the calculation means 205 calculates the FDD, the X-ray technician can easily determine the arrangement of the radiation source 101 and the two-dimensional detector 105. In particular, in a subject rotation type CBCT in which a subject rotates, which is about to be put into practical use, an X-ray technician has no imaging experience. For this reason, appropriately determining the FDD is an important factor that determines the quality of shooting.

  Next, the radiation source moving unit 206 arranges the focal point (which is also the position of the radiation source) 101 and the two-dimensional detector 105 at the FDD distance interval calculated by the calculating unit (S304). Here, a motor-driven moving mechanism (not shown) controlled by the control means 202 is used, but a general moving mechanism may be used. As a result, the X-ray technician can automatically obtain the arrangement of the focal point 101 and the two-dimensional detector 105 which are suitable distances by inputting the information input means 203. Thereby, there is an effect that accurate photographing can be performed quickly. This also has the effect of reducing the burden on the patient and the X-ray technician.

  Furthermore, when a mechanism for automatically operating the radiation source moving unit 206 cannot be provided, the focal point 101 and the two-dimensional image are artificially determined from the FDD that is the calculation result of the calculation unit 205 displayed by the display unit 211. The detector 105 may be arranged. In addition, when arrangement | positioning is performed artificially, FDD is confirmed with a potentiometer. Also in this case, there is an effect that it is possible to obtain the arrangement of the focal point 101 and the two-dimensional detector 105 which are suitable distances. Thereby, there is an effect that accurate photographing can be performed quickly. This also has the effect of reducing the burden on the patient and the X-ray technician.

  Next, an operation during photographing after the arrangement of the focal point 101 and the two-dimensional detector 105 will be described with reference to the flow of FIG.

  In accordance with the photographing start signal input from the interface unit 204, the control unit 202 transmits a rotation signal of the rotary table 102 to the rotary table motor control unit (S401). Then, the control unit 202 monitors an encoder signal (not shown) generated from the rotation table 102 that has started rotating, and confirms whether a predetermined constant speed and angle have been reached (S402). When a predetermined constant speed and angle are reached, the control means 202 sends a signal to the radiation source output means 208 to start X-ray exposure (S403) and data collection (S404). The encoder signal is also used to determine data integration timing. The subject 106 is rotated by the subject 106 standing on the rotary table 102.

  If an encoder that generates 25000 pulses per table rotation is used and 1000 views of projection data are collected per rotation, data is collected from the two-dimensional detector 105 every 25 pulses of the encoder signal. become. The control means 202 counts the encode pulses, generates an integrated signal every 25 pulses, and counts the X-ray dose reaching the two-dimensional detector 105. The photographing is continued until a predetermined count number is reached (S404, S405).

  In the present embodiment, it is assumed that X-rays are generated continuously, but the present invention is not limited to this, and pulses are generated in accordance with the integration interval of the two-dimensional detector 105 based on the encoder signal. X-rays may be generated. Data from the two-dimensional detector 105 is sequentially transferred to the reconstruction unit 210 via the BUS. The data transfer continues until the turntable 102 rotates a predetermined rotation angle and a predetermined number of views are collected (S406). When the rotary table 102 rotates a predetermined rotation angle and reaches a predetermined number of views, the control unit 202 instructs the output unit 208 of the radiation source to stop the X-ray exposure (S406). Thereafter, the rotary table 102 is controlled to be stopped while decelerating (S408).

  Immediately after the X-ray exposure is completed, the last projection data is transferred to the reconstruction unit 210. The control unit 202 instructs the reconstruction unit 210 to perform reconstruction based on the collected projection data. Reconfiguration may be started after the entire data collection is completed. The reconstruction includes preprocessing, filter processing, and back projection processing. The preprocessing includes offset processing, LOG conversion, gain correction, and defect correction. In the filter processing, a Ramachandran function or a Shepp Logan function is generally used, and these are also used in this embodiment. The filtered data is backprojected. The algorithm from the filtering process to the back projection uses the Feldkamp algorithm. When the back projection is completed and the CT cross-sectional image is reconstructed, the cross-section is displayed on the image display unit 211 (S409). The cross-sectional image is displayed and shooting is completed.

The reconstruction algorithm uses the Feldkamp algorithm, but is not limited thereto. References include a method described by Feldkamp, Davis and Kress ("Practical Cone-Beam Algorithm"), J. Opt. Soc. Am. 612-619, 1984. The geometric system is shown below.
Two-dimensional detector width 430 (mm)
Height in the Z direction of the two-dimensional detector 430 (mm)
Focus-rotation center distance FCD = 2223 (mm) (Focus-Center-Distance)
Focus-detector distance FDD = 2423 (mm) (Focus-Detector-Distance)
Reconstruction height HOV = 350 (mm) (Height of View)
Effective field diameter FOV = 389 (mm) (Field of View)
Cone angle φ = 10 degrees The height of the lung field is a statistically high range of about 35 cm. Based on this value, if the cone angle is 5 ° on one side (φ = 10 °), FCD = 2223 mm. . At the same time, if FOV = 389 mm and the gap between the outer edge of the FOV and the sensor is secured about 5 mm, FDD = 2423 mm. When the two-dimensional detector 105 is square, the actual condition is HOV = 389 mm, but a region exceeding 35 cm is considered to be a reconstructed image not suitable for diagnosis because the cone angle exceeds 5 ° on one side.

  Although it has been described above that the cone angle is preferably 5 ° or less on one side, experimental data for limiting the numerical value of 5 ° on one side is shown below. An experimental imaging system is shown in FIG. 5, and an experimental phantom placed on the rotary table 102 is placed at a location 900 mm from the tube. The experimental FPD is fixedly arranged at a distance of about 100 mm from the center of the experimental phantom. The resolution of the FPD used in the experiment is 0.64 × 0.64 mm, and the number of pixels is 384 pixels (horizontal) × 224 pixels (vertical).

  A cross-sectional view of the experimental phantom used is shown in FIG. The experimental phantom is a combination of acrylic plates in 6 directions, and the size of the cross section of the plate in the diameter direction is 100 mm. FIG. 7 shows an example of an actual reconstructed image. Since the number of reconstructed pixels is 384 × 384 pixels and the reconstructed area is 140 mmφ, the pixel size is 0.365 × 0.365 mm. The reconstruction algorithm uses the Feldkamp algorithm. In FIG. 6, a rectangular white portion indicates an ROI (Region Of Interest).

  In order to verify the limit value of the cone angle, the average pixel value and standard deviation of ROI shown in FIG. 6 are used as indices. FIG. 7 shows the average pixel value and standard deviation of ROI in each reconstructed cross section. The horizontal axis is the number of the reconstructed cross section. The cross section from 1 to 350 is reconstructed, and the reconstruction pitch of each reconstructed cross section is 0.365 mm, similar to the resolution of the cross section. That is, it is reconstructed in three dimensions with equal resolution. In the figure, series 1 is an average pixel value (lower waveform), and series 2 is a standard deviation (upper waveform).

  As can be seen from FIG. 7, it can be read that the standard deviation has deteriorated from the cross-sectional position A1. However, the poor standard deviation in the cross section 90-120 in the figure is due to sensor artifacts and not due to cone angle. Although the average pixel value and the standard deviation are worse after the cross-sectional position A2, this generally indicates a cross section where a straight line connecting the X-ray tube and the experimental FPD intersects the reconstruction area. That is, the deterioration of the average pixel value and standard deviation due to data loss is demonstrated.

  When the cone angle of the cross-sectional position A1 is obtained, it is 5.07 ° on one side. That is, the cone angle is 10.14 °. Next, the average value and standard deviation of the CT values of the air region in the reconstructed image are shown. The ROI of the air region is shown in FIG. 8, and the average value and standard deviation for each cross section are shown in FIG. It turns out that the tendency similar to an acrylic part is shown. However, in FIG. 9, the bad standard deviation in the cross section 90-120 is due to sensor artifacts and not due to cone angle.

  The results of evaluation by a plurality of doctors, etc., of images actually taken at a plurality of cone angles also support the above experimental results. Thus, the formula (1) is derived from a large number of experimental results and a large number of clinical evaluations, and is a formula for calculating a suitable FDD.

  If the cone angle is decreased, an image with less reconstruction error can be obtained. However, as the cone angle is decreased, the distance between the two-dimensional detector 105 and the radiation source 101 must be increased, and the cone angle is decreased. In this case, it is necessary to increase the X-ray dose output from the radiation source 101. However, in order to suppress the radiation dose of the radiation source 101 as much as possible and obtain a high-quality reconstructed image, it is desirable to install the radiation source 101 at a distance calculated by the equation (1).

  However, when there is a request to obtain a higher-quality reconstructed image, since the high-output radiation generation source 101 has been developed, it is possible to make FCD and FDD larger than the above numerical values. This makes it possible to obtain a suitable reconstructed image with a small reconstruction error.

Conventionally, in simple photographing, a film having a length of 43 cm and a width of 35 cm has been used. This is because even the person with the highest lung field has a limit of 43 cm. Therefore, if HOV = 430 mm is input to the above equation, FCD≈2730 mm. However, in the above formula (1), the reconstruction area FOV = 389 mm.phi., But when a general CT CT FOV = 500 mm.phi. Is adopted and the distance from the reconstruction area to the sensor is 100 mm, FDD = 250 mm + 100 mm + 0. 5 * HOV / tan5 (2)
When HOV = 430 mm is input to this, FDD≈3080 mm is obtained. Therefore, as a practical range of this radiation apparatus, 2400-3000 mm is a preferable range.

  Next, in a scene where the CBCT according to the present invention is actually used, it is necessary to photograph a large number of subjects in a short time, such as chest examination. In this case, it is also required to shoot with the FDD fixed for the majority of subjects. As described above, it is generally desirable to set the FDD to approximately 2400 mm in consideration of the majority of people with a lung field height of 35 cm or less and the suppression of the radiation dose of the radiation source as much as possible. Thereby, in the case of chest imaging, the movement time of the radiation source 101 can be eliminated, and the imaging efficiency is improved. Furthermore, the radiation dose of the radiation source can be reduced.

  In addition, considering a person with a lung field height of about 43 cm rarely, it is desirable to set the FDD in a range up to approximately 3000 mm. Also in this case, the movement time of the radiation source 101 can be eliminated, and there is an effect of increasing the imaging efficiency. Furthermore, the radiation dose of the radiation source can be reduced. Furthermore, there is an effect that it is possible to cope with photographing of a subject with a lung field height of 35 cm or more which is rarely photographed.

  Also, when considering operation in a large hospital, it is necessary to take a large amount of images of not only the chest but also the abdomen, abdomen, cervical spine, etc. Therefore, it is necessary to set the FDD to a distance at which an image with no image can be taken.

  In this case, as described above, it is desirable to set the FDD to 2400 mm or more that can support chest imaging. When the doctor prefers an image with less reconstruction error, it is desirable to set the FDD distance to 2400 mm or more within the range allowed by the output of the radiation source. In this way, when the FDD distance is set to 2400 mm or more, an image of a reconstruction error can be obtained, and there are effects that it is possible to deal with many imaging regions.

  Next, as described above, the higher-resolution reconstructed image can be obtained as the cone angle is reduced. However, the range of the cone angle will be examined below in view of practicality when the present CBCT is actually performed. This CBCT assumes that general imaging (which is imaging without reconstruction and is also referred to as simple imaging) can be performed when rotation is stopped. Therefore, the apparatus according to the present invention must be able to withstand imaging using a general X-ray apparatus. Note that the general X-ray apparatus for imaging can be produced at a lower cost than a tube for X-ray CT, and is widely used in the market from clinics to university hospitals.

The anode heat capacity of an X-ray tube used for X-ray CT is about 7000 KHU, whereas the tube anode heat capacity of a general radiography X-ray apparatus is 200-300 KHU. The cooling capacity per minute of the X-ray tube for general radiography is 15-20% of the anode heat capacity. This is due to the structure in which the radiation from the anode and the heat conduction from the bearing are received by the surrounding oil and the oil is air-cooled by the fan. Here, when the imaging dose of CBCT in the chest clinical practice is 120 [KV], N [mA], and 5 [seconds], the anode heat capacity HU by the imaging is empirically in the case of the inverter system,
120 [KV] X N [mA] X 5 [seconds] X 1.41 = 846 XN [HU] (3)
It is represented by

The following are conditions for clinical evaluation of the chest using the CBCT apparatus according to the present invention.
Two-dimensional detector width 430 (mm)
Height in the Z direction of the two-dimensional detector 430 (mm)
Focus-rotation center distance FCD = 3067 (mm) (Focus-Center-Distance)
Focus-detector distance FDD = 3417 (mm) (Focus-Detector-Distance)
Effective viewing area (FOV) FOV = 385mmφ
Cone angle φ = 7.2 degrees X-ray condition 120 [KV], 100 [mA]
X-ray filter Copper 1.5mm
Primary ray transmittance of grid for removing scattered radiation 30%
Scanning time per rotation: 5 seconds Under these conditions, image evaluation was performed by a doctor who conducted a chest clinical experiment on two human bodies. Although no disease was found, it was expected that a solid nodule of 10 mm could be reliably detected from the degree of blood vessel recognition that could be observed. A detection capability that can reliably detect a solid nodule of 10 mm is an effective device for medical examinations.

When the amount of X-ray heat at the time of imaging under the above conditions is obtained from Equation (3), it is 84.6 KHU. Although it is empirically obvious that the image quality of the apparatus according to the present invention can be improved if the X-ray dose is increased, as described above, if the heat capacity of the tube is taken into consideration, increasing the X-ray dose is a patient. As exposure increases, the shooting interval is extended. Therefore, if the image provided under the imaging conditions is used as a reference, the relationship between the X-ray heat quantity, the tube cooling capacity C [KHU / min], and the imaging interval T [min] satisfies Expression (4). The standard image quality can be obtained.
C [KHU / min] * T [min] ≧ 84.6 [KHU] (4)
Equation (4) is applied to a system using a general X-ray apparatus. The tube cooling capacity C [KHU / min] is C = 60 [KHU / min] when the anode capacity is 300 KHU and the cooling capacity is 20%. It can be seen that Equation (4) is satisfied when the shooting interval is 2 minutes. Here, the shorter one of the average intervals of 2-3 minutes performed at the examination site was adopted as the imaging interval. When the maximum FDD (FDDm) is calculated under these conditions, FDDm ≦ 4069 mm is obtained from the equation (5).
120KHU ≧ 84.6KHU * (FDDm / 3417) 2 (5)
In summary, when the anode capacity D [KHU] or less, the cooling capacity E [1 / min] or less, the imaging interval T [min] or less, the reconstruction height (HOV), and the effective visual field (FOV), The FDD lower limit (FDDmin) and the FDD upper limit (FDDmax) can be obtained by the following equations.
FDDmin = 0.5 * FOV + 0.5 * HOV / tan (φ / 2) Equation (6)
FDDmax = 3417 * SQRT (D * E * T / 84.6) Formula (7)
Here, φ represents the cone angle, and the maximum is 10 degrees by experiment. Substituting HOV = 350 mm and FOV = 389 mm into equation (6) yields FDD = 2418 mm. In CBCT using a general imaging apparatus, it can be seen that in order to realize an imaging interval of 2 minutes, the FDD must be set to approximately 4 m or less. When the FDD is 4 m, the cone angle φ = 6 degrees.

It is a figure which shows the structural example of CBCT of a to-be-photographed object type | mold whose subject rotates. 1 is a diagram illustrating a system configuration example of a subject rotation type CBCT in which a subject rotates. FIG. It is a flowchart which shows the flow of the process which determines FDD. It is a flowchart which shows the flow of imaging | photography of CBCT. It is a figure which shows the imaging | photography form used for experiment. It is sectional drawing of an experiment phantom. It is a figure which shows the average pixel value of ROI, and a standard deviation. It is sectional drawing of an experiment phantom. It is a figure which shows the average pixel value of ROI, and a standard deviation.

Explanation of symbols

DESCRIPTION OF SYMBOLS 101 Radiation source 102 Rotary table 105 Two-dimensional detector 106 Subject 203 Information input means 205 Calculation means 206 Radiation source movement means 210 Reconstruction means 211 Image display means

Claims (5)

  1. A radiation source for exposing the subject to radiation,
    A rotating means for relatively rotating a subject in the radiation emitted by the radiation source; a two-dimensional detector for detecting the radiation;
    Reconstructing means for reconstructing an output signal from the two-dimensional detector,
    When imaging while rotating the subject relatively by the rotating means, the output signal is reconstructed as CT imaging,
    The radiation imaging apparatus according to claim 1, wherein the imaging is not reconfigured as general imaging when imaging is performed with the rotation stopped.
  2.   2. The apparatus according to claim 1, wherein when imaging is performed while the subject is relatively rotated by the rotating means, the radiation source and the two-dimensional detector are arranged at an interval of 240 cm or more and 400 cm or less. Radiography equipment.
  3.   The radiographic apparatus according to claim 1, wherein the two-dimensional detector is 43 cm or more in the rotation axis direction.
  4.   2. The radiation imaging apparatus according to claim 1, wherein the radiation source has a tube anode heat capacity of 300 KHU or less.
  5. The shooting height (HOV) is 35 cm or more, the effective field diameter (FOV) is 39 cm or more, the tube anode heat capacity of the radiation source is 300 KHU or less, the tube cooling capacity is 20% / min or less, and the shooting interval is 2 minutes or less. In this case, the radiation imaging apparatus according to claim 1, wherein the radiation source and the two-dimensional detector are arranged at an interval of 240 cm or more and 400 cm or less.
JP2006144354A 2003-06-09 2006-05-24 Radiation imaging apparatus Pending JP2006218327A (en)

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Cited By (2)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
JP2011058983A (en) * 2009-09-11 2011-03-24 Hitachi Ltd Method for photographing of radiation tomograph
JP2011125486A (en) * 2009-12-17 2011-06-30 Toshiba Corp X-ray ct system and control method for same

Cited By (2)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
JP2011058983A (en) * 2009-09-11 2011-03-24 Hitachi Ltd Method for photographing of radiation tomograph
JP2011125486A (en) * 2009-12-17 2011-06-30 Toshiba Corp X-ray ct system and control method for same

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