JP5298806B2 - Tomography equipment - Google Patents

Description

  The present invention relates to a tomography apparatus used for a CT apparatus, a C-type arm apparatus, and the like.

  FIG. 7 shows a C-type arm device in cone beam reconstruction. X-rays emitted from the X-ray tube 101 are detected by an X-ray detector 102 typified by a flat panel X-ray detector (FPD). In the middle, the X-rays pass through the subject M placed on the top (bed) and attenuate, and the X-rays are scattered by the subject M and the top, and scattered rays are also generated. The X-rays attenuated together with the scattered rays pass through the X-ray grid installed in front of the X-ray detector 102, thereby reducing the scattered rays. In CT data collection, a projection image is collected over a half circumference while rotating the C-arm 104.

  In the reconstruction processing for creating a tomographic image, the acquired projection image data is converted into transmission length data of the subject and processed, and then a tomographic image is obtained. When calculating transmission length data by the subject from the collected projection image data, calculate the signal attenuation amount from the projection signal image collected in advance without the subject (hereinafter referred to as pre-attenuation signal data). Thus, the transmission length is obtained. Normally, the signal data before attenuation is obtained by averaging a sufficient number of frames for noise reduction.

  In cone beam CT, since many tomographic planes are reconstructed at a time, the influence of scattered radiation is greater than in a normal CT apparatus. The X-ray grid described above is arranged with a gap with respect to a large number of thin plate-like X-ray absorbing foils (grid foils), which reduces many scattered rays. The direct line is also somewhat blocked. In particular, when the grid foil is arranged in the horizontal direction or the vertical direction on the projection image, the shadow of the grid foil appears clearly on the projection image depending on the condition of the pixel interval of the X-ray detector and the grid foil interval. .

  Normally, if there is a shadow on the grid foil, there is a problem in diagnosing from the obtained image, so a process for removing the shadow is performed. However, in practice, the shadow of the grid foil that appears to be removed on the projection image but cannot be removed by the reconstruction processing of the tomographic image may affect the tomographic image.

  In particular, in the C-type arm device, a CT device (a X-ray tube and an X-ray detector are arranged in a gantry surrounding the body axis of the subject, Compared to a CT device that rotates around the body axis of the subject in the gantry, the X-ray tube and the X-ray detector are held at both ends of the C-type arm. It is easily affected by arm deflection. Specifically, the relative position of the X-ray tube with respect to the X-ray detector surface changes according to the weight in accordance with the position of the rotating C-arm (that is, the collection angle position). Therefore, the shadow position of the grid foil also changes according to the collection angle position, and the signal before attenuation when nothing passes through the subject changes according to the collection angle position. However, if these changes are ignored and the image reconstruction processing is performed by averaging the signal data before attenuation, a false image is generated and the image quality of the tomographic image (reconstructed image) is degraded.

  In order to prevent generation of a false image, signal data before attenuation is collected for each collection angle position while rotating the C-arm even when there is no subject, similarly to CT data collection when the subject is present. Using these collected pre-attenuation signal data for each collection angle position as a reference, the data is converted into attenuation and transmission water permeation length (water length conversion), and a tomographic image (reconstruction) is performed. Image) (see FIG. 3). In this method, since the image reconstruction is performed in consideration of the deflection of the C-arm and the change caused by the deflection, false images generated by the deflection can be reduced.

  In general, the sensitivity characteristic that is the ratio of the detected value to the input X-ray is not always the same in all detection pixels on the X-ray detector. The thickness of the detection substance, X-ray energy, amplification factor, etc. are related. Therefore, a projection image (usually projection data without an object) under a condition where the incident X-rays are considered to be constant is acquired in advance, and correction is performed so that the variation in data at this time becomes uniform. Here, this correction processing is referred to as “sensitivity correction”.

  Ideally, X-ray detectors have the same sensitivity characteristics regardless of the signal range. However, it is divided into a high signal region and a low signal region with a certain intermediate signal region (hereinafter referred to as "detector specific signal region") as a boundary, and different sensitivity characteristics in the high signal region and the low signal region. If it has, there may be a case where there are a plurality of such borders, or the sensitivity characteristics may gradually change without a border clearly.

  For example, when using X-ray detectors having different sensitivity characteristics in the high signal region and the low signal region with the detector specific signal region as a boundary, the signal data before attenuation for each collection angle position described above is used. The following problems occur when seeking to reduce the CT data. That is, the pre-attenuation signal data has a signal value that does not pass through the subject (that is, a high signal value before the attenuation) and is in a high signal range, so sensitivity correction in the high signal range must be performed. On the other hand, the signal value that has passed through the subject (that is, the low signal value after attenuation) is a low signal range, and sensitivity correction must be performed in the low signal range. The correction does not work well.

  Where sensitivity must be corrected in the low signal range, because the attenuation was calculated without considering the difference in sensitivity characteristics, the subject region above a certain thickness would be a signal lower than the detector specific signal range, In particular, a false image is generated in the central portion of the reconstructed image having a certain thickness or more. Therefore, there is a drawback in that the image quality deteriorates particularly at a place where it is desired to examine the subject.

  In order to solve this problem, when collecting the pre-attenuation signal data for each collection angle position described above, a method of irradiating X-rays with a reduced dose and collecting the detection signal in a low signal region is conceivable. In this case, since sensitivity correction is performed at the low signal region level, appropriate sensitivity correction is performed for CT data having the same low signal region. Therefore, it is possible to avoid the disadvantage that the image quality is deteriorated due to generation of a false image at a location where the subject is particularly desired to be examined (location other than the surface layer portion). Sensitivity correction does not work well on the surface layer of the subject, but since there are few regions of interest in the surface layer, it is not a problem. However, when collecting data in the low signal range, the signal becomes smaller and the noise increases relatively. By using this data to calculate the amount of attenuation (attenuated data), the image quality deteriorates rather than the effect of noise. There is a fear.

  To relatively reduce this increasing noise, there is a method of increasing the number of data to be collected. However, for example, the method of slowly rotating the C-arm and collecting a lot of data during that time has a different rotation speed of the C-arm compared to the CT data collection at the normal rotation speed. Unlike the phenomenon that occurs in CT data collection, it may not be used for correction. Further, in order to rotate the C-arm in the same way as CT data collection at a normal rotation speed, data collection must be performed as many times as necessary, and the operation time until data collection is completed is long. It takes time and effort.

Furthermore, a method of obtaining attenuation data after correcting the linearity of the detector sensitivity characteristic and performing CT reconstruction is known (for example, see Patent Document 1).
JP 2006-26412 A

  However, in the case of the method as described in Patent Document 1 described above, data acquisition for correcting linearity requires various data collection for each tube voltage condition and the entire acquired signal, which is troublesome. .

  The present invention has been made in view of such circumstances, and an object of the present invention is to provide a tomographic apparatus capable of suppressing deterioration in image quality caused by two or more different signal areas.

In order to achieve such an object, the present invention has the following configuration.
In other words, the invention described in claim 1 is a tomographic apparatus that acquires a tomographic image by tomography, and uses a difference in projection value distribution from each projection value in two or more different signal areas as a distribution difference physical quantity. Using the distribution difference physical quantity calculation means to be calculated and the distribution difference physical quantity obtained by the distribution difference physical quantity calculation means, attenuation data or projection data affected by the sensitivity characteristics in one signal range is converted to sensitivity characteristics in another signal range. And a data conversion means for converting the data into the data affected by the above, and obtaining a tomographic image based on the data converted by the data conversion means.

  [Operation / Effect] According to the invention described in claim 1, the distribution difference physical quantity calculating means obtains the distribution difference of the projection values from each projection value in two or more different signal areas as the distribution difference physical quantity. Using the distribution difference physical quantity obtained by the distribution difference physical quantity calculation means, the attenuation data or projection data affected by the sensitivity characteristic in one signal range is converted into data affected by the sensitivity characteristic in another signal range. The conversion means converts. Therefore, even when two or more different signal regions have different sensitivity characteristics, image quality degradation caused by two or more different signal regions can be obtained only by obtaining a projection value in the corresponding signal region. Can be suppressed.

  When two or more different signal regions are divided into a high signal region and a low signal region with a detector specific signal region as a boundary, in one example of the above-described invention, the distribution difference physical quantity calculating means is as follows: Then, the distribution difference physical quantity is obtained, and the data conversion means converts the data as follows. That is, the distribution difference physical quantity calculation means converts the distribution of the projection value affected by the sensitivity characteristic in the low signal area from the two different signal areas to the distribution of the projection value affected by the sensitivity characteristic in the high signal area. To calculate the distribution difference physical quantity. Using the distribution difference physical quantity obtained by the distribution difference physical quantity calculation means, data attenuation or projection data affected by the sensitivity characteristics in the low signal range is converted to data affected by the sensitivity characteristics in the high signal range Means convert. Therefore, even when there are different sensitivity characteristics between the high signal area and the low signal area with the detector specific signal area as the boundary, it is only necessary to obtain projection values in the two signal areas. It is possible to suppress deterioration in image quality due to the area. For example, it is possible to suppress deterioration in image quality at portions other than the surface layer of the subject.

  According to the tomographic apparatus according to the present invention, even when two or more different signal regions have different sensitivity characteristics, two or more different ones can be obtained only by acquiring projection values in the corresponding signal regions. Degradation of image quality due to the signal range can be suppressed.

Embodiments of the present invention will be described below with reference to the drawings.
FIG. 1 is a schematic diagram and a block diagram of a C-type arm device according to the embodiment, FIG. 2 is a block diagram in which the flow of each data in the C-type arm device according to the embodiment is shown, and FIG. FIG. 4 is a block diagram showing a flow of each conventional data for comparison with FIG. 2, and FIG. 4 is a flowchart showing a series of processes according to the embodiment.

  As shown in FIG. 1, the C-type arm device according to this embodiment detects an X-ray tube 1 that irradiates X-rays, and an X-ray that is irradiated from the X-ray tube 1 and passes through a subject M or the like. C-type holding an X-ray tube 2 and an X-ray detector 2, an arithmetic processing unit 3 that performs various arithmetic processes based on X-ray detection signals detected by the X-ray detector 2 Arm 4 is provided. In FIG. 1, the illustration of the X-ray grid is omitted. The X-ray tube 1 and the X-ray detector 2 are held at both ends of the C-type arm 4, and the C-type arm 4 rotates around the subject M, whereby the X-type held by the C-type arm 4 respectively. The tube 1 and the X-ray detector 2 rotate around the subject M.

The X-ray detector 2 converts X-rays irradiated from the X-ray tube 1 and passing through the subject M and entering the detection surface into charge signals, and further converts the charge signals into electric signals (detection signals). X-rays are detected by collecting the values. In the present embodiment, the detection signal without the subject in the high signal range detected with the C-arm 4 stationary is collected from the X-ray detector 2 as a high signal range projection image High [i]. Similarly, the detection signal without the subject in the low signal region detected in a state where the C-arm 4 is stationary is collected from the X-ray detector 2 as the low signal region projection image Low [i]. In addition, the detection signal without the subject detected while rotating the C-arm 4 in the same manner as collecting the CT data in the state without the subject M is the pre-attenuation signal data P ₀ [i] [j ] From the X-ray detector 2 and the detection signal with the subject M detected while rotating the C-arm 4 is collected from the X-ray detector 2 as normal CT data CT [i] [j] To do. Note that i represents an index of the detection pixel. When the X-ray detector 2 is configured by M detection pixels, i = 1, 2, 3,... j represents an index of projection angles, and j = 1, 2, 3,..., J when collected at J projection angles.

  The arithmetic processing unit 3 includes a central processing unit (CPU). The arithmetic processing unit 3 includes a distribution difference parameter calculation unit 31, an attenuation amount calculation unit 32, a data conversion unit 33, a water length conversion unit 34, and a reconstruction processing unit 35. As will be described later with reference to the conventional block diagram in FIG. 3, the present embodiment is characterized in that the arithmetic processing unit 3 newly includes a distribution difference parameter calculation unit 31 and a data conversion unit 33.

  The distribution difference parameter calculation unit 31 calculates the difference in projection value distribution from the high signal area projection image High [i] and the low signal area projection image Low [i] collected by the X-ray detector 2 described above. Calculate as [i]. That is, the distribution difference parameter LtoH [i] for converting the distribution of the projection value affected by the sensitivity characteristic in the low signal range to the distribution of the projection value affected by the sensitivity characteristic in the high signal range is obtained. The distribution difference parameter corresponds to the distribution difference physical quantity in the present invention.

The attenuation amount calculation unit 32 obtains an attenuation amount (attenuation data) that is a detection signal ratio of X-rays before and after transmission. In the present embodiment, the CT data CT [i] [, which is the detection signal after transmission, is the pre-attenuation signal data P ₀ [i] [j] which is the detection signal before transmission collected by the X-ray detector 2 described above. j]. If attenuation data is Att [i] [j], it is represented by CT [i] [j] / P ₀ [i] [j].

  The data conversion unit 33 converts the attenuation data or the projection data affected by the sensitivity correction in the low signal range into the data affected by the sensitivity correction in the high signal range. In this embodiment, the attenuation data Att [i] [j] affected by the sensitivity correction in the low signal range obtained by the attenuation amount calculation unit 32 is used as the attenuation data Att affected by the sensitivity correction in the high signal range. The case of converting to [i] [j] 'will be described as an example.

  The water length conversion unit 34 obtains the attenuation data Att [i] [j] ′ converted by the data conversion unit 33 by converting it into the permeation length of water, which is a permeation material required for tomography. The reconstruction processing unit 35 performs a reconstruction process based on the transmission length of water converted by the water length conversion unit 34 and acquires a tomographic image (reconstructed image).

  The distribution difference parameter calculation unit 31 corresponds to the distribution difference physical quantity calculation unit in the present invention, and the data conversion unit 33 corresponds to the data conversion unit in the present invention. Specific functions of each component in the arithmetic processing unit 3 will be described later with reference to a flowchart of FIG.

  Next, specific functions of the components in the arithmetic processing unit 3 will be described with reference to FIGS. For comparison with FIG. 2, the conventional data flow will be described with reference to FIG.

(Step S1) Collecting with the C-arm stationary The X-ray detector 1 irradiates X-rays from the X-ray tube 1 while the C-arm 4 is stationary without placing the subject on the top. To do. When collecting a detection signal without a subject in a high signal range, X-rays are irradiated from the X-ray tube 1 at a tube current and irradiation time that are about the same as those of normal CT data collection. When collecting detection signals without a subject in a low signal range, X-rays with a reduced dose are irradiated. In this embodiment, X-rays are irradiated from the X-ray tube 1 for a plurality of times (for example, 30 frames / second for 1 second), and the detection signals for the plurality of times are collected by the X-ray detector 2 and the average value thereof is collected. Are collected as projection images. For example, the detection signal without the subject in the high signal range is collected, and the average value of the detection signal without the subject in the high signal range for 30 frames is set as the high signal range projection image High [i], and the low signal The detection signals without the subject in the region are collected, and the average value of the detection signals without the subject in the low signal region for 30 frames is defined as the low signal region projection image Low [i]. As shown in FIG. 2, the high signal area projection image High [i] and the low signal area projection image Low [i] are sent to the distribution difference parameter calculation unit 31.

  First, in order to collect a detection signal without a subject in a high signal region, X-rays are continuously irradiated from the X-ray tube 1 and then a detection signal without a subject in a low signal region is collected. Therefore, the X-ray tube 1 may be continuously irradiated with X-rays, and conversely, the X-ray tube 1 may be continuously irradiated with X-rays in order to collect detection signals without a subject in a low signal range. After the irradiation, X-rays may be continuously irradiated from the X-ray tube 1 in order to collect detection signals without a subject in a high signal range. Further, in order to alternately collect the detection signal without the subject in the high signal region and the detection signal without the subject in the low signal region, the X-ray tube 1 is continuously irradiated with X-rays alternately. Also good. Moreover, although the average value of the detection signal without the subject in the high / low signal region for a plurality of times is the projection image, it is not limited to the average value, and may be a median value or the like. Therefore, a median value of detection signals without an object in a plurality of high / low signal regions may be used as a projection image.

(Step S2) Distribution Difference Parameter Calculation The distribution difference parameter calculation unit 31 distributes projection values from the high signal area projection image High [i] and the low signal area projection image Low [i] collected by the X-ray detector 2. Is obtained as a distribution difference parameter LtoH [i]. Specifically, the distribution difference parameter LtoH [i] is expressed by the following equation (1).

LtoH [i] = (High [i] / Low [i]) * average (Low [i] / High [i]) (1)
Here, average (x [i]) represents a spatial average value of x [i]. Therefore, average (x [i]) = Σx [i] / M (the sum of x [1] to x [M] divided by the number M of detected pixels) represents average (x [i]). The As shown in FIG. 2, the distribution difference parameter LtoH [i] is sent to the data converter 33.

(Step S3) Collection by rotating the C-arm without a subject and irradiating X-rays from the X-ray tube 1 while rotating the C-arm 4 without placing the subject on the top plate X-ray detector 2 Collect with. By rotating the C-arm 4 around the subject M, the X-ray detector 2 while the X-ray tube 1 and the X-ray detector 2 respectively held by the C-arm 4 rotate around the subject M. Sequentially collects detection signals without a subject. The detection signal without the subject is the pre-attenuation signal data P ₀ [i] [j] for each collection angle position. As shown in FIG. 2, the pre-attenuation signal data P ₀ [i] [j] is sent to the attenuation amount calculation unit 32. For the purpose of noise reduction, there are cases in which P ₀ [i] [j] are bundled in the j direction or averaged.

(Step S4) Collection by rotating the C-arm with the subject Normal CT data collection is performed. That is, the subject M is placed on the top and the X-ray detector 1 emits X-rays while the C-arm 4 is rotated and is collected by the X-ray detector 2. By rotating the C-arm 4 around the subject M, the X-ray detector 2 while the X-ray tube 1 and the X-ray detector 2 respectively held by the C-arm 4 rotate around the subject M. Sequentially collects detection signals with the subject M. The detection signal with the subject M is normal CT data CT [i] [j] for each collection angle position. As shown in FIG. 2, this CT data CT [i] [j] is sent to the attenuation amount calculation unit 32.

(Step S5) Calculation of attenuation data The attenuation amount calculation unit 32 uses the detection signal after transmission as the detection signal before transmission collected by the X-ray detector 2 as the detection signal before transmission P ₀ [i] [j]. A certain CT data CT [i] [j] is divided, and the value obtained by the division, that is, CT [i] [j] / P ₀ [i] [j] is attenuated data Att [i] [j] Asking. As shown in FIG. 2, the attenuation data Att [i] [j] is sent to the data converter 33.

(Step S6) Data Conversion The data conversion unit 33 uses the distribution difference parameter LtoH [i] obtained by the distribution difference parameter calculation unit 31 to use the attenuation data Att [i] [j obtained by the attenuation amount calculation unit 32. ] Is converted into attenuation data Att [i] [j] '. The attenuated data Att [i] [j] ′ after conversion is expressed by the following equation (2).

Att [i] [j] '= Att [i] [j] * LtoH [i] (2)
It is converted into attenuation data Att [i] [j] ′ that is made uniform in the low signal range by the above equation (2). As shown in FIG. 2, the attenuated data Att [i] [j] ′ after conversion is sent to the water length conversion unit 34.

  As apparent from the description of FIG. 2, conventionally, as shown in FIG. 3, when the attenuation amount calculation unit 132 obtains the attenuation data Att [i] [j], the attenuation data Att [i] [j ] Is sent to the water length conversion unit 134, in this embodiment, when the attenuation amount calculation unit 32 obtains the attenuation data Att [i] [j] as shown in FIG. Att [i] [j] is sent to the data conversion unit 33, and the data conversion unit 33 converts it into attenuation data Att [i] [j] ′ that is made uniform in the low signal range. The converted attenuation data Att [i] [j] ′ is sent to the water length conversion unit 34. Therefore, conventionally, there is no concept of distribution difference parameter LtoH [i] used for data conversion and data conversion. In other words, in the present embodiment, a distribution difference parameter calculation unit 31 and a data conversion unit 33 that are not conventionally provided are newly provided.

(Step S7) Water Length Conversion / Reconstruction Process The water length conversion unit 34 converts the attenuation data Att [i] [j] ′ converted by the data conversion unit 33 into a water permeation length and obtains it. That is, the attenuation data Att [i] [j] ′ is further corrected based on the known linear attenuation coefficient of water, and the transmission length associated with the corrected attenuation data is obtained. Since specific water length conversion is known, its description is omitted. Based on the converted water permeation length, the reconstruction processing unit 35 solves the inverse problem, performs reconstruction processing, and obtains the distribution of the permeable material, thereby obtaining a tomographic image (reconstructed image). Since the specific reconstruction process is known, the description thereof is omitted.

  According to the above-described C-type arm device according to the present embodiment, the distribution difference parameter calculation unit 31 uses each projection value (this book) in two or more different signal regions (in this embodiment, a high signal region and a low signal region). In the embodiment, a difference in projection value distribution is obtained as a distribution difference parameter from the high signal area projection image High [i] and the low signal area projection image Low [i]). Using the distribution difference parameter obtained by the distribution difference parameter calculation unit 31, attenuation data Att [i] [j] affected by the sensitivity characteristic in a certain signal range (low signal range in this embodiment) The data conversion unit 33 converts the data into attenuation data Att [i] [j] ′ influenced by the sensitivity characteristic in the signal range (high signal range in this embodiment). Therefore, even when two or more different signal regions (high / low signal regions) have different sensitivity characteristics, the projection value (high signal region projection image High [i] and low signal in the corresponding signal region). The image quality degradation due to different sensitivity characteristics in two or more different signal regions (high / low signal regions) can be suppressed only by acquiring the region projection image Low [i]).

  As in the present embodiment, when two or more different signal regions are divided into a high signal region and a low signal region with a detector specific signal region as a boundary, the distribution difference parameter calculation unit 31 mutually A distribution difference parameter LtoH [i] for converting the projection value distribution in the low signal range to the projection value distribution in the high signal range among the two different signal ranges is obtained. Using the distribution difference parameter LtoH [i] obtained by the distribution difference parameter calculation unit 31, the attenuation data Att [i] [j] affected by the sensitivity characteristic in the low signal range is converted into the sensitivity characteristic in the high signal range. The data conversion unit 33 converts the attenuation data Att [i] [j] ′ affected by Therefore, even when the high signal region and the low signal region have different sensitivity characteristics with the detector specific signal region as a boundary, the high signal region projection image High [i] indicating the projection values in the two signal regions and By only acquiring the low signal area projection image Low [i], it is possible to suppress the deterioration of the image quality due to the difference in sensitivity characteristics between the high signal area and the low signal area.

  By converting to attenuation data Att [i] [j] ′ that is made uniform in the low signal range, the signal at the time of transmission through the subject M becomes accurate, and other than the surface layer portion of the subject M (for example, the region of interest) The false image at the center portion of the subject M is reduced. On the other hand, although a false image is generated in the surface layer portion of the subject M, there is no problem because the region of interest is the central portion of the subject M. In this way, it is possible to suppress deterioration in image quality at portions other than the surface layer portion of the subject M.

  The present invention is not limited to the above-described embodiment, and can be modified as follows.

  (1) In the above-described embodiment, all data is converted into data affected by the sensitivity characteristics in the high signal range. Conversely, from the distribution of the projection values affected by the sensitivity characteristics in the high signal range. A distribution difference parameter (distribution difference physical quantity) for converting to a distribution of projected values affected by the sensitivity characteristics in the low signal range is obtained, and attenuation data affected by the sensitivity characteristics in the high signal range is obtained using the distribution difference parameters. Alternatively, the projection data may be converted into data affected by sensitivity characteristics in the low signal range.

  (2) In the above-described embodiment, the case of two different signal regions (high / low signal region in the embodiment) has been described as an example. However, if two or more different signal regions are targets, The present invention may be applied to, for example, a case where there are a plurality of boundaries and there are three or more signal areas different from each other. Therefore, there may be two boundaries, and there may be three signal regions, ie, a high signal region, a medium signal region, and a low signal region.

(3) In the above-described embodiment, the attenuation data Att [i] [j] affected by the sensitivity characteristic in the low signal range is changed to the attenuation data Att [i] [j] affected by the sensitivity characteristic in the high signal range. In the above example, the conversion target is the attenuation data. However, as shown in FIG. 6, the conversion target may be projection data. That is, after converting projection data of a certain signal range (low signal range in FIG. 6) into projection data affected by sensitivity characteristics in another signal range (high signal range in FIG. 6), the projection data is used. You may ask for attenuation data. In FIG. 6, the projection data is the pre-attenuation signal data P ₀ [i] [j], and the pre-attenuation signal data P ₀ [i] [j] is attenuated using the distribution difference parameter LtoH [i]. The data conversion unit 33 converts the previous signal data P ₀ [i] [j] ′ into the previous signal data P ₀ [i] [j] ′, and converts the pre-attenuation signal data P ₀ [i] [j] ′ after the conversion into the pre-attenuation signal data P ₀ [i] [ Dividing by j], the attenuation amount calculation unit 32 obtains attenuation data Att [i] [j]. Since the obtained attenuation data Att [i] [j] has already been converted, it may be sent directly to the water length converter 34.

  (4) In the above-described embodiments, the tomography apparatus is a C-type arm apparatus, but the present invention may be applied to an X-ray CT apparatus that is a dedicated machine. Thus, the tomographic apparatus to which the present invention is applied is not particularly limited.

It is the schematic and block diagram of a C-type arm apparatus which concern on an Example. It is the block diagram which wrote together the flow of each data in the C-type arm apparatus which concerns on an Example. FIG. 3 is a block diagram showing a flow of each conventional data for comparison with FIG. 2. It is a flowchart which shows a series of processes which concern on an Example. It is the figure which showed typically the case where a gain changes gradually between a low signal area and a high signal area. It is the block diagram which wrote together the flow of each data in the C-type arm apparatus which concerns on a modification. It is the schematic of a C-type arm apparatus.

Explanation of symbols

31 ... Distribution difference parameter calculation unit 33 ... Data conversion unit
High [i] ... High signal area projection image
Low [i] ... Low signal projection image
LtoH [i]… Distribution difference parameter
P ₀ [i] [j]… Signal data before attenuation
CT [i] [j] ... CT data
Att [i] [j]… Attenuation data
Att [i] [j] '… attenuated data after conversion

Claims (2)

  A tomography apparatus for obtaining a tomographic image by tomography, a distribution difference physical quantity calculation means for obtaining a distribution difference of projection values from each projection value in two or more different signal areas as a distribution difference physical quantity, and the distribution difference Data conversion that converts attenuation data or projection data affected by sensitivity characteristics in one signal range into data affected by sensitivity characteristics in another signal range using the distribution difference physical quantity obtained by the physical quantity calculation means Means for obtaining a tomographic image based on the data converted by the data conversion means.   The tomography apparatus according to claim 1, wherein the distribution difference physical quantity calculating unit is configured to detect sensitivity in a high signal region from a distribution of projection values affected by sensitivity characteristics in a low signal region among two different signal regions. The distribution difference physical quantity to be converted into the distribution of the projection value affected by the characteristic is obtained, and the distribution difference physical quantity obtained by the distribution difference physical quantity calculating means is used to receive the influence of the sensitivity characteristic in the low signal range. The tomography apparatus, wherein the data conversion means converts the attenuated data or projection data into data affected by sensitivity characteristics in the high signal range.

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