JP6985004B2 - Photon counting type X-ray CT device and image processing device - Google Patents

Description

Embodiments of the present invention relate to a photon counting type X-ray CT apparatus and an image processing apparatus.

The photon counting type X-ray CT apparatus detects the number and energy of X-ray photons incident on the detection element, generates a CT image for each different energy, discriminates substances, and estimates the concentration of substances. This is also called multi-energy reconstruction.

However, when a large number of photons are incident on the detection element at one time, the photon counting type X-ray CT apparatus may not be able to properly discriminate substances and estimate the concentration of substances. This is because the waveforms generated by the individual photons overlap due to the response time of the detection element, and a plurality of photons are counted as one photon. This phenomenon is called pulse pile-up.

Japanese Unexamined Patent Publication No. 2000-23965 Japanese Unexamined Patent Publication No. 2007-167663

An object to be solved by the present invention is to provide a photon counting type X-ray CT apparatus and an image processing apparatus capable of appropriately performing a diagnosis using a reconstructed image.

The photon counting type X-ray CT apparatus according to the embodiment includes an X-ray tube, a detector, a photon counting unit, a correction unit, and a calculation unit. The X-ray tube irradiates the subject with X-rays. The detector has a plurality of detection elements for detecting the incident X-ray photons. The photon counting unit counts the number of photons of the X-ray for each energy band set on the energy distribution of the X-ray, the position of the X-ray tube, and the detection element. The correction unit corrects the number of X-ray photons counted by the photon counting unit based on the detection characteristics of the detection element. The calculation unit calculates the reliability of the pixels of the reconstructed image according to the correction.

FIG. 1 is a diagram showing a configuration example of a photon counting type X-ray CT apparatus according to an embodiment. FIG. 2 is a flowchart showing an example of processing performed by the photon counting type X-ray CT apparatus according to the embodiment. FIG. 3 is a diagram showing an example of a subject having a uniform X-ray attenuation coefficient. FIG. 4 is a diagram showing the relationship between the incident angle of X-rays transmitted through points A, B, C, D and E in FIG. 3 and the transmission distance of X-rays transmitted through these points. FIG. 5 is a diagram showing the relationship between the incident angle of X-rays transmitted through points A, B, C, D and E in FIG. 3 and the transmitted dose of X-rays transmitted through these points. FIG. 6 is a diagram showing the relationship between the dose of X-rays incident on the detection element and the number of photons of X-rays. FIG. 7 is a diagram showing the relationship between the dose of X-rays incident on the detection element and the reciprocal of the error generated when the correction function corrects the number of photons of X-rays. FIG. 8 is a diagram showing an example of reliability displayed together with a morphological image of a subject.

Hereinafter, the photon counting type X-ray CT apparatus and the image processing apparatus according to the embodiment will be described with reference to the drawings. In the following embodiments, duplicate description will be omitted as appropriate.

(Embodiment)
The configuration of the photon counting type X-ray CT apparatus 1 according to the embodiment will be described with reference to FIG. FIG. 1 is a diagram showing a configuration example of a photon counting type X-ray CT apparatus according to an embodiment. As shown in FIG. 1, the photon counting type X-ray CT apparatus 1 includes a gantry 10, a sleeper 20, and a console 30. The configuration of the photon counting type X-ray CT apparatus 1 is not limited to the following configuration.

The gantry 10 includes a high voltage generation circuit 11, a collimator adjustment circuit 12, a gantry drive circuit 13, an X-ray irradiation device 14, a detector 15, a data acquisition circuit 16, and a rotating frame 17.

The high voltage generation circuit 11 supplies a tube voltage to the X-ray tube 141, which will be described later. The high voltage generation circuit 11 realizes its function by reading and executing a program stored in the storage circuit 35 described later.

The collimator adjusting circuit 12 adjusts the opening degree and the position of the collimator 143, which will be described later. As a result, the collimator adjustment circuit 12 adjusts the irradiation range of the X-rays that the X-ray tube 141 irradiates the subject P. The collimator adjustment circuit 12 realizes its function by reading and executing a program stored in a storage circuit 35 described later.

The gantry drive circuit 13 rotates and drives the rotary frame 17. As a result, the gantry drive circuit 13 turns the X-ray irradiation device 14 and the detector 15 in a circular orbit centered on the subject P. The gantry drive circuit 13 realizes its function by reading and executing a program stored in a storage circuit 35, which will be described later.

The X-ray irradiation device 14 includes an X-ray tube 141, a wedge 142, and a collimator 143. The X-ray tube 141 irradiates the subject P with X-rays. The X-ray tube 141 generates beam-shaped X-rays by the tube voltage supplied by the high voltage generation circuit 11. This beam-shaped X-ray is also called a cone beam. The wedge 142 is an X-ray filter for adjusting the dose of X-rays emitted from the X-ray tube 141. The collimator 143 is a slit for adjusting the irradiation range of X-rays. The opening degree and position of the collimator 143 are adjusted by the collimator adjusting circuit 12. By adjusting the opening degree of the collimator 143, for example, the fan angle and the cone angle of the cone beam are adjusted.

The detector 15 has a plurality of detection elements for detecting incident X-ray photons. The plurality of detection elements are regularly arranged in the first direction and the second direction intersecting the first direction. For example, the first direction is the channel direction and the second direction is the slice direction. Here, the channel direction is the circumferential direction of the rotating frame 17, and the slice direction is the Z direction. In addition, such a detector is called a multi-row detector.

The detection element includes a scintillator, a photodiode, a photon counting type detection circuit, and an integral type detection circuit. A detector in which the detection element has a scintillator and a photodiode is called a solid-state detector. The input terminal of the photon counting type detection circuit is connected to the output terminal of the photodiode. The output terminal of the photon counting type detection circuit is connected to the input terminal of the data acquisition circuit 16. The input terminal of the integrator detection circuit is connected to the output terminal of the photodiode. The output terminal of the integral type detection circuit is connected to the input terminal of the data acquisition circuit 16.

The detection element converts the incident X-ray photons into a response waveform or a voltage pulse by the following method. The detection element converts the incident X-ray photons into light by a scintillator. The detection element converts this light into an electric charge by a photodiode. This charge is output to the photon counting type detection circuit and the integral type detection circuit. The photon counting type detection circuit outputs a response waveform to the data acquisition circuit 16. The response waveform is time-series data of the voltage generated by each photon incident on each detection element at each position of the X-ray tube 141. The integral type detection circuit outputs a voltage pulse to the data acquisition circuit 16. The voltage pulse is the sum of the voltages generated by all the photons incident on each detection element at each position of the X-ray tube 141.

The detector 15 may be a direct conversion type detector. The direct conversion type detector is a detector that directly converts X-rays incident on a detection element into electric charges. This charge is at least that the electrons generated by the incident of X-rays move toward the positive potential current collecting electrode and the holes generated by the incident of X-rays move toward the negative potential current collecting electrode. Output by one.

The data acquisition circuit 16 has a photon counting function 161, a dosimetry function 162, a projection data generation function 163, and a conversion function 164. Details of these functions will be described later. The data acquisition circuit 16 is realized by, for example, a processor. The data acquisition circuit 16 is also called a DAS (Data Acquisition System).

The rotating frame 17 is an annular frame that supports the X-ray irradiation device 14 and the detector 15 so as to face each other with the subject P in between. The rotating frame 17 is driven by the gantry drive circuit 13 and rotates at high speed on a circular orbit centered on the subject P.

The sleeper 20 includes a top plate 21 and a sleeper drive circuit 22. The top plate 21 is a plate-shaped member on which the subject P is placed. The sleeper drive circuit 22 moves the subject P in the shooting port of the gantry 10 by moving the top plate 21 on which the subject P is placed in the body axis direction. The sleeper drive circuit 22 realizes its function by reading and executing a program stored in the storage circuit 35 described later.

The console 30 includes an input circuit 31, a display 32, a projection data storage circuit 33, an image storage circuit 34, a storage circuit 35, and a processing circuit 36.

The input circuit 31 is used by a user who inputs an instruction or a setting. The input circuit 31 is included in, for example, a mouse and a keyboard. The input circuit 31 transfers the instructions and settings input by the user to the processing circuit 36. The input circuit 31 is realized by, for example, a processor.

The display 32 is a monitor referenced by the user. The display 32 is, for example, a liquid crystal display or an organic EL (Electroluminescence) display. The display 32 receives, for example, an instruction from the processing circuit 36 to display a reconstructed image and a GUI (Graphical User Interface) used when a user inputs an instruction or a setting. The display 32 displays, for example, a reconstructed image and a GUI based on this instruction. The reconstructed image is, for example, a morphological image, a substance discrimination image, an electron density image, an effective atomic number image, and a monochromatic X-ray image. Morphological images are generated by reconstructing integral projection data. Alternatively, the morphological image is generated by reconstructing photon counting type projection data obtained by adding up the number of X-ray photons in a plurality of energy bands. The photon counting type X-ray CT apparatus 1 does not have to have the display 32.

The projection data storage circuit 33 stores raw data (Raw Data) generated by the preprocessing function 362 described later. The image storage circuit 34 stores the reconstructed image generated by the image generation function 363 described later.

The storage circuit 35 stores a program for the high voltage generation circuit 11, the collimator adjustment circuit 12, the gantry drive circuit 13, and the data acquisition circuit 16 to realize the above-mentioned functions. The storage circuit 35 stores a program for the sleeper drive circuit 22 to realize the above-mentioned functions. The storage circuit 35 stores a program for the processing circuit 36 to realize each of the functions described later.

The projection data storage circuit 33, the image storage circuit 34, and the storage circuit 35 have a storage medium in which the stored information can be read out by a computer. The storage medium is, for example, a hard disk.

The processing circuit 36 has a scan control function 361, a preprocessing function 362, an image generation function 363, a correction function 364, a calculation function 365, a display control function 366, and a control function 367. The correction function 364 corrects the number of X-ray photons counted by the photon counting function 161 based on the detection characteristics of the detection element. The image generation function 363 generates a reconstructed image based on the result corrected by the correction function 364. The calculation function 365 calculates the reliability of the pixels of the reconstructed image based on the result corrected by the correction function 364. Details of these functions will be described later. The processing circuit 36 is realized by, for example, a processor.

An example of the processing of the photon counting type X-ray CT apparatus 1 according to the embodiment will be described with reference to FIGS. 2 to 8. FIG. 2 is a flowchart showing an example of processing performed by the photon counting type X-ray CT apparatus according to the embodiment. Further, in the description of an example of the processing performed by the photon counting type X-ray CT apparatus 1, FIGS. 3 to 8 are appropriately referred to.

As shown in FIG. 2, the processing circuit 36 scans the subject, counts the number of X-ray photons, and measures the X-ray dose (step S1). The process of step S1 is, for example, as follows.

The processing circuit 36 reads a program corresponding to the scan control function 361 from the storage circuit 35 and executes it. The scan control function 361 is a function of controlling the photon counting type X-ray CT apparatus 1 in order to execute a scan. The processing circuit 36 controls the photon counting type X-ray CT apparatus 1 as follows, for example, by executing the scan control function 361.

The processing circuit 36 controls the sleeper drive circuit 22 to move the subject P into the photographing port of the gantry 10. The processing circuit 36 controls the gantry drive circuit 13 to scan the subject P. Specifically, the processing circuit 36 controls the high voltage generation circuit 11 to supply the tube voltage to the X-ray tube 141. The processing circuit 36 adjusts the opening degree and the position of the collimator 143 by controlling the collimator adjusting circuit 12. Further, the processing circuit 36 rotates the rotating frame 17 by controlling the gantry drive circuit 13.

The scans performed by the photon counting type X-ray CT apparatus 1 are, for example, a conventional scan, a helical scan, and a step-and-shoot. The conventional scan is a method of scanning the subject P with the position of the subject P placed on the top plate 21 fixed. Helical scan is a method of scanning a subject P while moving the subject P placed on the top plate 21 in the body axis direction. The step-and-shoot is a method in which the position of the subject P placed on the top plate 21 is moved at regular intervals to perform conventional scanning in a plurality of scan areas.

The processing circuit 36 controls the data acquisition circuit 16 to count the number of X-ray photons and measure the X-ray dose. The data acquisition circuit 16 performs the following processing, for example.

The data acquisition circuit 16 reads a program corresponding to the photon counting function 161 from the storage circuit 35 and executes it. The photon counting function 161 counts the number of continuous events in which the response waveform output by the photon counting type detection circuit exceeds a predetermined threshold value. As a result, the photon counting function 161 counts the number of X-ray photons incident on the detection element. Further, the photon counting function 161 calculates the wave height and the waveform area of the response waveform output by the detection element. As a result, the photon counting function 161 calculates the energy of the X-ray photons incident on the detection element. Therefore, the photon counting function 161 can count the number of X-ray photons for each energy band set on the X-ray energy distribution, the position of the X-ray tube 141, and the detection element.

The data acquisition circuit 16 reads a program corresponding to the dosimetry function 162 from the storage circuit 35 and executes it. The dosimetry function 162 integrates the voltage pulse output by the integral type detection circuit. As a result, the dosimetry function 162 measures the dose of X-rays incident on the detection element.

As shown in FIG. 2, the processing circuit 36 generates photon counting type projection data and integral type projection data (step S2). The process of step S2 is, for example, as follows.

The data acquisition circuit 16 reads a program corresponding to the projection data generation function 163 from the storage circuit 35 and executes it. The projection data generation function 163 is a function of generating projection data based on the response waveform or voltage pulse output by the detection element. The projection data is, for example, a synogram. The synogram is data in which the number of X-ray photons incident on each detection element or the dose of X-rays is arranged at each position of the X-ray tube 141. Here, the position of the X-ray tube 141 is called a view. That is, the synogram is data in which the number of X-ray photons or the dose of X-rays is assigned to a two-dimensional Cartesian coordinate system centered on the view direction and the channel direction. The projection data generation function 163 generates a synogram for each column in the slice direction, for example.

The projection data generation function 163 generates photon counting type projection data based on the number of X-ray photons counted by the photon counting function 161. The brightness of each pixel of the photon counting type projection data represents the energy band set on the energy distribution of the X-ray, the position of the X-ray tube, and the number of X-ray photons counted for each detection element. The projection data generation function 163 generates integral projection data based on the X-ray dose measured by the dosimetry function 162. The brightness of each pixel of the integral projection data represents the dose of X-rays incident on each position of the X-ray tube 141 and each detection element.

The data acquisition circuit 16 may read and execute the program corresponding to the conversion function 164 from the storage circuit 35 before reading and executing the program corresponding to the projection data generation function 163 from the storage circuit 35. The conversion function 164 calculates the attenuation coefficient of X-rays when passing through the subject P from the number of X-ray photons counted by the photon counting function 161 or from the number of X-ray photons counted by the photon counting function 161. This is a function for calculating the transmission distance of X-rays when transmitting through the subject P.

In this case, the projection data generation function 163 determines the X-ray attenuation coefficient when passing through the subject P or the X-rays when passing through the subject P, based on the number of X-ray photons counted by the photon counting function 161. Generates photon counting type projection data that represents the transmission distance. The energy distribution (energy spectrum) of the number of photons counted at the coordinates on a certain photon counting type projection data is I, and the energy spectrum of the X-ray before being exposed from the X-ray tube 141 and incident on the subject P is I ₀ . Assuming that the energy is E and the attenuation coefficient of the X-ray incident on the coordinates of the subject P is μ, the following equation (1) holds.

Here, L represents the transmission distance of the subject P. If the subject P is made of one kind of substance and the attenuation coefficient μ is known, the equation (1) can be converted into a transmission distance by the following equation (2).

However, since the subject P is usually a patient, it is composed of a plurality of substances. Further, in the case of contrast imaging, the X-ray CT apparatus 1 is required to generate a reconstructed image in which the contrast agent is discriminated. Therefore, for example, assuming two basal substances (two of muscle tissue, bone, contrast medium, etc.), the formula (2) becomes the following formula (3).

The conversion function 164 calculates _{the permeation distances L 1} and L _{2 of the two substances.} The permeation distances L ₁ and L ₂ of the two substances are calculated in a plurality with respect to the energy E. Therefore, the conversion function 164 statistically solves them to obtain the transmission distances L ₁ and L ₂ . The number of basal substances is not particularly limited.

Further, when the transmission distance L (= L ₁ + L ₂ ) is obtained in advance from the integral type projection data, the conversion function 164 can obtain the attenuation coefficient. If there is an assumption of the basal material, the conversion function 164 can also convert from the mass attenuation coefficient of the basal material to the density of the material.

The projection data generation function 163 can generate photon counting type projection data of the transmission distance and the attenuation coefficient by using the transmission distance and the attenuation coefficient calculated in this way as the luminance values of the pixels of the projection data.

As shown in FIG. 2, the processing circuit 36 performs preprocessing on the integral projection data (step S3). The process of step S3 is, for example, as follows.

The processing circuit 36 reads a program corresponding to the preprocessing function 362 from the storage circuit 35 and executes it. The preprocessing function 362 is a function for correcting the projection data generated by the data acquisition circuit 16. This correction is, for example, logarithmic transformation, offset correction, sensitivity correction, beam hardening correction, and scattered ray correction. The integral projection data corrected by the preprocessing function 362 is stored in the projection data storage circuit 33. The integral projection data corrected by the preprocessing function 362 is also referred to as raw data.

As shown in FIG. 2, the processing circuit 36 reconstructs the integral projection data and generates a morphological image (step S4). The process of step S4 is, for example, as follows.

The processing circuit 36 reads a program corresponding to the image generation function 363 from the storage circuit 35 and executes it. The image generation function 363 includes a function of reconstructing the integral type projection data stored in the projection data storage circuit 33 and generating a morphological image. The reconstruction method is, for example, a back projection process or a successive approximation method. Further, the back projection process is, for example, an FBP (Filtered Back Projection) method. The morphological image generated by the image generation function 363 is stored in the image storage circuit 34. The morphological image generated in step S4 may be either a two-dimensional morphological image or a three-dimensional morphological image.

The processing circuit 36 calculates the transmitted dose of X-rays transmitted through each point of the subject based on the morphological image (step S5). The process of step S5 is, for example, as follows.

The processing circuit 36 reads a program corresponding to the correction function 364 from the storage circuit 35 and executes it. The correction function 364 calculates the transmitted dose of X-rays transmitted through the points of the morphological image generated in step S4 corresponding to the points of the subject P and incident on the detection element. Specifically, the correction function 364 transmits the points based on the value calculated by line-integrating the CT values with a straight line passing through each point of the morphological image generated in step S4 corresponding to each point of the subject P. Calculate the transmitted dose of X-rays. Alternatively, the correction function 364 may calculate the transmitted dose of X-rays transmitted through the points at each point of the subject P based on the brightness of each pixel of the integral projection data. The X-ray transmission dose calculated in step S5 depends on the incident angle of the X-ray. Therefore, with reference to FIGS. 3, 4 and 5, the behavior of the transmitted dose of X-rays calculated in step S5 with respect to the incident angle of X-rays will be described.

FIG. 3 is a diagram showing an example of a subject having a uniform X-ray attenuation coefficient. In the description of an example of the processing performed by the photon counting type X-ray CT apparatus 1, the case where the subject P1 shown in FIG. 3 is scanned will be given as an example. The subject P1 is an elliptical pillar whose central axis is parallel to the Z axis and whose bottom surface is parallel to the XY plane. Further, the X-ray attenuation coefficient of the subject P1 is uniform.

FIG. 4 is a diagram showing the relationship between the incident angle of X-rays transmitted through points A, B, C, D and E in FIG. 3 and the transmission distance of X-rays transmitted through these points. In the following description, it is assumed that the X-rays passing through these five points pass through a plane parallel to the XY plane. Further, the incident angle of the X-ray transmitted through these points is defined by the angle θ measured from the half straight line extending in the + X direction from the point A.

As shown in FIG. 4, the behavior of the transmission distance of the X-ray transmitted through the point A with respect to the incident angle is a downwardly convex curve symmetrical with respect to the point where the incident angle is 90 degrees. Further, as shown in FIG. 4, the transmission distance of the X-ray transmitted through the point A is maximum when the incident angle is 0 degrees and 180 degrees, and is minimum when the incident angle is 90 degrees. The maximum value of the transmission distance of X-rays transmitted through the point A is equal to the length of the long axis of the bottom surface of the subject P1. The minimum value of the transmission distance of X-rays transmitted through the point A is equal to the length of the short axis of the bottom surface of the subject P1. These are due to the fact that the point A is located on the central axis of the subject P1.

FIG. 5 is a diagram showing the relationship between the incident angle of X-rays transmitted through points A, B, C, D and E in FIG. 3 and the transmitted dose of X-rays transmitted through these points. In FIG. 5, the dose of X-rays when the transmission distance of the subject P1 is zero is set to 100. The behavior of the dose of X-rays transmitted through points A, B, C, D and E shown in FIG. 3 with respect to the incident angle is, as shown in FIG. 5, above and below each curve shown in FIG. The curve looks like the opposite. This is because the line attenuation coefficient of the subject P1 is uniform.

As shown in FIG. 4, the behavior of the transmission distance of the X-ray transmitted through the point B, the point C, or the point D with respect to the incident angle is a downwardly convex curve. The maximum value of the transmission distance of X-rays transmitted through the point B, the point C, or the point D is smaller than the length of the long axis of the bottom surface of the subject P1. The maximum value of the transmission distance of X-rays transmitted through the point B, the point C, or the point D is smaller than the length of the short axis of the bottom surface of the subject P1. Further, the minimum value of the transmission distance of the X-rays transmitted through the point C is smaller than the minimum value of the transmission distance of the X-rays transmitted through the point B, and larger than the minimum value of the transmission distance of the X-rays transmitted through the point C. Further, the incident angle at which the transmission distance of the X-rays transmitted through the point C is minimized is smaller than the incident angle at which the transmission distance of the X-rays transmitted through the point B is minimized, and is the transmission distance of the X-rays transmitted through the point C. Greater than the minimum. In these, the points B, C and D are located in the second quadrant of the XY coordinates centered on the point on the central axis of the subject P1, and the point C is closer to the surface of the subject P1 than the point B. This is because it is located farther from the surface of the subject P1 than the point D.

As shown in FIG. 4, the behavior of the transmission distance of the X-ray transmitted through the point E with respect to the incident angle includes a portion that decreases monotonically, a portion that is constant at zero, and a portion that increases monotonically. Further, the incident angle at which the transmission distance of the X-ray transmitted through the point E becomes zero is smaller than 90 degrees. These are the second quadrants of the XY coordinates centered on the point on the central axis of the subject P1, and the point E is located in the region outside the subject P1.

As described above, the dose of X-rays transmitted through each point of the subject P1 and incident on the detection element differs depending on the incident angle even at the same point. Therefore, even at the same point of the subject P1, the influence of the pulse pile-up on the number of X-ray photons counted by the photon counting function 161 differs depending on the incident angle of the X-rays.

The processing circuit 36 corrects the photon counting type projection data based on the transmitted dose calculated in step S5 (step S6). The process of step S6 is, for example, as follows.

The processing circuit 36 reads a program corresponding to the correction function 364 from the storage circuit 35 and executes it. The correction function 364 corrects the number of X-ray photons counted by the photon counting function 161 for each energy band set on the X-ray energy distribution, the position of the X-ray tube 141, and the detection element based on the calculated dose. do. That is, in the correction function 364, the number of X-ray photons counted by the photon counting function 161 is proportional to the number of X-ray photons incident on the detection element by the number of X-ray photons counted by the photon counting function 161. Correct to the assumed number. The correction function 364 uses the relationship shown in FIG. 6 when performing this correction.

FIG. 6 is a diagram showing the relationship between the dose of X-rays incident on the detection element and the number of photons of X-rays. The dotted line D shown in FIG. 6 shows the relationship between the two when it is assumed that the number of X-ray photons counted by the photon counting function 161 is proportional to the dose of X-rays incident on the detection element. The solid line S shown in FIG. 6 represents the behavior of the number of X-ray photons counted by the photon counting function 161 with respect to the dose of X-rays incident on the detection element. That is, the solid line S represents the detection characteristic of the detection element. As shown in FIG. 6, as the dose of X-rays incident on the detection element increases, the difference between the number of X-ray photons indicated by the dotted line D and the number of X-ray photons indicated by the solid line S increases. .. This is because the pulse pile-up is more likely to occur as the X-ray dose increases.

The correction function 364 multiplies the number of X-ray photons shown by the solid line S by a predetermined constant at each dose, and corrects the number of X-ray photons shown by the dotted line D. This predetermined constant depends on the dose. That is, the solid line S represents the behavior of the number of X-ray photons before correction with respect to the X-ray dose. The dotted line D represents the behavior of the corrected number of X-ray photons with respect to the X-ray dose.

The error bar on the solid line S represents an error in each dose of the number of X-ray photons counted by the photon counting function 161. These errors are, for example, in the range of twice the standard deviation of the number of X-ray photons counted by the photon counting function 161 at each dose. The error bar on the dotted line D represents an error in each dose of the number of X-ray photons corrected by the correction function 364. These errors are, for example, a range obtained by multiplying the error of the number of X-ray photons counted by the photon counting function 161 at each dose by the above-mentioned predetermined constant. As shown in FIG. 6, the larger the dose of X-rays incident on the detection element, the larger the error in each dose of the number of X-ray photons corrected by the correction function 364. The above-mentioned error may be calculated based on the Poisson distribution of the number of X-ray photons incident on the detection element.

As shown in FIG. 2, the processing circuit 36 performs preprocessing on the corrected photon counting type projection data (step S7). Similar to step S3, the processing circuit 36 reads out a program corresponding to the preprocessing function 362 from the storage circuit 35 and executes it. The preprocessing function 362 applies the same correction as in step S3 to the photon counting type projection data. The photon counting type projection data corrected by the preprocessing function 362 is stored in the projection data storage circuit 33. The photon counting type projection data corrected by the preprocessing function 362 is also called raw data.

As shown in FIG. 2, the processing circuit 36 reconstructs the preprocessed photon counting type projection data to generate a substance discrimination image (step S8). Similar to step S4, the processing circuit 36 reads out a program corresponding to the image generation function 363 from the storage circuit 35 and executes it. The image generation function 363 includes a function of reconstructing the photon counting type projection data stored in the projection data storage circuit 33 and generating a substance discrimination image. The reconstruction method is as described in step S4. The substance discrimination image generated by the image generation function 363 is stored in the image storage circuit 34. The image generation function 363 may generate at least one of a substance discrimination image, an electron density image, an effective atomic number image, and a monochromatic X-ray image.

As shown in FIG. 2, the processing circuit 36 calculates the reliability based on the result of the correction performed in step S6 (step S9). The process of step S9 is, for example, as follows.

The processing circuit 36 reads a program corresponding to the calculation function 365 from the storage circuit 35 and executes it. The calculation function 365 is a function of calculating the reliability of the pixels of the reconstructed image based on the result corrected by the correction function 364. The calculation function 365 calculates the reliability as follows, for example.

FIG. 7 is a diagram showing the relationship between the dose of X-rays incident on the detection element and the reciprocal of the error generated when the correction function corrects the number of photons of X-rays. The threshold value Th shown in FIG. 7 determines whether or not the number of X-ray photons corrected by the correction function 364 is reliable. The threshold value Th is determined by, for example, the calculation function 365. If the reciprocal of the error exceeds the threshold Th, the number of X-ray photons corrected by the correction function 364 is determined to be reliable. When the reciprocal of the error is equal to or less than the threshold value Th, the number of X-ray photons corrected by the correction function 364 is determined to be unreliable.

As shown in FIG. 7, the reciprocal of the error exceeds the threshold Th in the range of the dose of X-rays incident on the detection element in the range of 0 to 60. Therefore, when the dose of X-rays incident on the detection element is in the range of 0 to 60, the number of X-ray photons corrected by the correction function 364 is judged to be reliable. As shown in FIG. 7, the reciprocal of the error is a threshold value Th or less in the range where the dose of X-rays incident on the detection element is in the range of 60 to 100. Therefore, when the dose of X-rays incident on the detection element is in the range of 60 to 100, the number of X-ray photons corrected by the correction function 364 is determined to be unreliable.

The calculation function 365 calculates the reliability at each point of the subject P1 based on the error generated when the correction function 364 corrects the number of X-ray photons transmitted through the point. For example, at each point of the subject P1, the calculation function 365 covers an angle range in which the reciprocal of the error generated when the correction function 364 corrects the number of X-ray photons passing through the point exceeds a predetermined threshold value. The value divided by is calculated as the reliability.

First, the calculation function 365 specifies an angle range in which the reciprocal of the error generated when correcting the number of X-ray photons passing through the point at each point of the subject P1 exceeds the threshold value Th. That is, the calculation function 365 specifies an angle range in which the dose of X-rays transmitted through the points in the subject P1 is in the range of 0 to 60. Next, the calculation function 365 calculates the value obtained by dividing the angle range in which the dose is in the range of 0 to 60 by the entire angle range as the reliability. When there are a plurality of angle ranges in which the dose is in the range of 0 to 60, the calculation function 365 calculates the value obtained by dividing the total angle range by the entire angle range as the reliability. When the scan control function 361 performs a full scan, the entire angle range means 360 degrees. Further, when the scan control function 361 performs a half scan, the entire angle range means 180 degrees.

As shown in FIG. 2, the processing circuit 36 displays a morphological image, a substance discrimination image, and a reliability (step S10). The process of step S10 is, for example, as follows.

The processing circuit 36 reads a program corresponding to the display control function 366 from the storage circuit 35 and executes it. The display control function 366 is a function for displaying the reconstructed image and the reliability on the display 32.

The display control function 366 displays the reliability together with the reconstructed image, for example, as shown in FIG. Specifically, for example, the display control function 366 causes the display 32 to display the reliability by the closed curve L80, the closed curve L90, and the closed curve L100. The closed curve L80 indicates that the reliability of the luminance of each pixel in the region surrounded by the closed curve is 80 or more. The closed curve L90 indicates that the reliability of the luminance of each pixel in the region surrounded by the closed curve is 90 or more. The closed curve L100 represents that the reliability of the luminance of each pixel in the region surrounded by the closed curve is 100.

Further, as shown in FIG. 8, the display control function 366 displays the reliability on the substance discrimination image IM. Specifically, the display control function 366 displays the closed curve L80, the closed curve L90, and the closed curve L100 indicating the reliability on the substance discrimination image IM. Further, the display control function 366 may display the reliability when the displayed cursor moves, for example. Specifically, the display control function 366 may display the closed curve L80, the closed curve L90, and the closed curve L100 indicating the reliability when the displayed cursor moves. The display control function 366 may display the reliability on the electron density image, the effective atomic number image, or the monochromatic X-ray image instead of the substance discrimination image IM.

The photon counting type X-ray CT apparatus 1 according to the embodiment has been described above. As described above, the processing circuit 36 corrects the number of X-ray photons counted by the photon counting function 161 based on the detection characteristics of the detection element by the correction function 364, and the correction function 364 corrects by the calculation function 365. The reliability of the pixels of the reconstructed image is calculated based on the above, and the reliability is displayed on the display 32. Therefore, the photon counting type X-ray CT apparatus 1 can provide the image interpreter with a reconstructed image capable of easily performing an appropriate diagnosis.

As described with reference to FIGS. 3, 4 and 5, the closer the point through which the X-ray is transmitted to the surface of the subject, the shorter the transmission distance of the X-ray transmitted through the subject, and the X incident on the detection element. The dose of the line increases. This phenomenon is particularly likely to occur in the photon counting type X-ray CT apparatus 1. This is because the photon counting type X-ray CT apparatus 1 irradiates the subject with X-rays from various directions in order to collect projection data. Therefore, the above-mentioned effect is important for the photon counting type X-ray CT apparatus 1.

Even if the brightness of each pixel of the photon counting type projection data represents the attenuation coefficient or transmission distance calculated by the conversion function 164, the calculation function 365 performs the same processing as in the case of correcting the number of photons, and is reliable. It is possible to calculate the degree.

The photon counting type X-ray CT apparatus 1 does not have to measure the X-ray dose. That is, the detection element does not have to have an integral type detection circuit. In this case, the data acquisition circuit 16 reads out a program corresponding to the dosimetry function 162 from the storage circuit 35 and executes it. The dosimetry function 162 calculates the dose of X-rays incident on the detection element by, for example, energy integrating the number of X-ray photons counted by the photon counting function 161.

The scan control function 361 has a dose higher than that of the X-rays previously irradiated to the X-ray tube 141 when the statistic that is an index of the reliability of the pixels included in the region of interest of the reconstructed image is smaller than a predetermined threshold. The subject P may be scanned by irradiating with a low X-ray. Here, the statistic that is an index of the magnitude of reliability is not a statistic that indicates the spread of reliability such as standard deviation and variance, but the minimum value, maximum value, median value, average value, and other statistics of reliability. It means a statistic that represents the size. The region of interest may also be set over the entire reconstructed image.

The scan control function 361 is, for example, a statistic that is an index of the reliability of pixels included in the region of interest of the reconstructed image generated from the projection data collected by the X-ray of the dose 80 shown in FIG. When is smaller than a predetermined threshold value, the X-ray tube 141 may be irradiated with X-rays having a dose of 40 shown in FIG. 6 to scan the subject P. As a result, the photon counting type X-ray CT apparatus 1 can generate a reconstructed image that accurately depicts the region of interest.

However, in this case, the scan control function 361 does not lower the X-ray dose to the level at which statistical fluctuations occur. This is due to the following reasons. If the dose of X-rays incident on the detection element is too low, the charge output by the detection element to the photon counting type detection circuit and the integral type detection circuit becomes too small. Therefore, the projection data generation function 163 cannot generate projection data having a high signal-to-noise ratio.

The correction function 364 can perform the process of step S5 even when the X-rays passing through each point of the subject pass through the plane intersecting the XY plane. Also in this case, the correction function 364 calculates the distance between the two points where the X-ray and the surface of the subject intersect as the transmission distance. Also in this case, the transmission distance depends on the incident angle of the X-ray.

The calculation function 365 may calculate a value used for diagnosis and associate this value with the reliability. The calculation function 365 applies computer-aided diagnosis (CAD) to the reconstructed image, for example, and calculates coordinates and volumes indicating a range to which radiotherapy is applied. The calculation function 365 associates the coordinates and volume indicating the range of radiation therapy with the reliability. Further, the display control function 366 may display on the display 32 the coordinates and the volume indicating the range to be subjected to the radiotherapy and the reliability associated with the coordinates and the volume. In this case, the method of displaying the reliability is not particularly limited. Alternatively, the display control function 366 may display an error in coordinates or volume indicating a range to be subjected to radiotherapy calculated based on the associated reliability. In this case, the error display method is not particularly limited.

Alternatively, the calculation function 365 calculates the stenosis rate of blood vessels based on the substance discrimination image of calcium. The calculation function 365 associates the stenosis rate with the reliability. Further, the display control function 366 may display the stenosis rate and the reliability associated with the stenosis rate on the display 32. In this case, the method of displaying the reliability is not particularly limited. Alternatively, the display control function 366 may display the error of the stenosis rate calculated based on the associated reliability. In this case, the error display method is not particularly limited.

Confidence is an indicator of the reliability of the values used for diagnosis. Therefore, the reliability is a judgment material when making a diagnosis using this value. Confidence is, for example, a judgment factor when planning radiation therapy.

If the appropriate radiation therapy plan changes depending on the coordinates and volume of the radiation therapy range, the doctor will change the radiation therapy plan in consideration of the radiation therapy range and its reliability. For example, if there is a substance that significantly attenuates the radiation used in radiation therapy between the surface of the subject and the area where radiation therapy is applied, the doctor considers the area where radiation therapy is applied and its reliability, and the radiation passes through. Change the area to be used. Alternatively, the doctor excludes the range of radiation therapy whose reliability is less than a predetermined threshold from the target of radiation therapy. In this case, the display control function 366 may display an image on the display 32 indicating that there is a range in which the reliability is smaller than a predetermined threshold value in the range to which the radiation therapy is applied. Physicians may also use electron density images to plan radiation therapy. In this case, the doctor formulates a radiation therapy plan in consideration of the range and dose distribution of radiation in the subject P calculated by the calculation function 365 based on the electron density and the reliability associated with these. ..

The calculation function 365 may calculate a value other than the value described in step S9 as the reliability. The calculation function 365 can calculate, for example, the values described below as the reliability. In the following description, the case where the reciprocal of the error behaves as shown in FIG. 7 with respect to the dose of X-rays incident on the detection element and the threshold value Th is determined as shown in FIG. 7 will be taken as an example.

The calculation function 365 sets the reciprocal of the error in an angle range in which the reciprocal of the error generated when the correction function 364 corrects the number of X-ray photons passing through the point at each point of the subject P exceeds a predetermined threshold value. Is calculated as the reliability by dividing by the angle range where is equal to or less than the predetermined threshold.

First, the calculation function 365 specifies an angle range in which the reciprocal of the error generated when correcting the number of X-ray photons passing through the point at each point of the subject P exceeds the threshold value Th. That is, the calculation function 365 specifies an angle range in which the dose of X-rays transmitted through the points in the subject P is in the range of 0 to 60. Next, the calculation function 365 specifies an angle range in which the reciprocal of the error generated when correcting the number of X-ray photons passing through the point at each point of the subject P is equal to or less than the threshold value Th. That is, the calculation function 365 specifies an angle range in which the dose of X-rays transmitted through the points in the subject P is in the range of 60 to 100.

Then, the calculation function 365 calculates the value obtained by dividing the angle range in which the dose is in the range of 0 to 60 by the angle range in which the dose is in the range of 60 to 100 as the reliability. When there are a plurality of angle ranges in which the dose is in the range of 0 to 60, the calculation function 365 calculates the reliability using the total angle range. Further, when there are a plurality of angle ranges in which the dose is in the range of 60 to 100, the calculation function 365 calculates the reliability using the total angle range.

Alternatively, the calculation function 365 integrates the reciprocal of the error that occurs when the correction function 364 corrects the number of X-ray photons passing through the point at each point of the subject P within the range of the incident angle of the X-ray. Thereby, the reliability of the pixel is calculated. The calculation function 365 is, for example, the reciprocal of the error in an angle range in which the reciprocal of the error generated when the correction function 364 corrects the number of X-ray photons passing through the point at each point of the subject P exceeds a predetermined threshold. The value obtained by dividing the value obtained by integrating the above by the value obtained by integrating the reciprocal of the error over the entire angle range is calculated as the reliability.

First, the calculation function 365 specifies an angle range in which the reciprocal of the error generated when correcting the number of X-ray photons passing through the point at each point of the subject P exceeds the threshold value Th. That is, the calculation function 365 specifies an angle range in which the dose of X-rays transmitted through the points in the subject P is in the range of 0 to 60. Next, the calculation function 365 calculates the value obtained by dividing the value obtained by integrating the reciprocal of the error in the angle range in which the dose is in the range of 0 to 60 by the value obtained by integrating the reciprocal of the error in the entire angle range as the reliability. ..

Alternatively, the calculation function 365 integrates the reciprocal of the error that occurs when the correction function 364 corrects the number of X-ray photons passing through the point at each point of the subject P within the range of the incident angle of the X-ray. Thereby, the reliability of the pixel is calculated. The calculation function 365 is, for example, the reciprocal of the error in an angle range in which the reciprocal of the error generated when the correction function 364 corrects the number of X-ray photons passing through the point at each point of the subject P exceeds a predetermined threshold. The value obtained by dividing the reciprocal of the error by the integrated value of the reciprocal of the error in an angle range in which the reciprocal of the error is equal to or less than a predetermined threshold value is calculated as the reliability.

First, the calculation function 365 specifies an angle range in which the reciprocal of the error generated when correcting the number of X-ray photons passing through the point at each point of the subject P exceeds the threshold value Th. That is, the calculation function 365 specifies an angle range in which the dose of X-rays transmitted through the points in the subject P is in the range of 0 to 60. Next, the calculation function 365 specifies an angle range in which the reciprocal of the error generated when correcting the number of X-ray photons passing through the point at each point of the subject P is equal to or less than the threshold value Th. That is, the calculation function 365 specifies an angle range in which the dose of X-rays transmitted through the points in the subject P is in the range of 60 to 100.

Then, the calculation function 365 divides the value obtained by integrating the reciprocal of the error in the angle range in which the dose is in the range of 0 to 60 by the value obtained by integrating the reciprocal of the error in the angle range in which the dose is in the range of 60 to 100. The value divided by the value is calculated as the reliability. When there are a plurality of angle ranges in which the dose is in the range of 0 to 60, the calculation function 365 calculates the reliability using the total angle range. Further, when there are a plurality of angle ranges in which the dose is in the range of 60 to 100, the calculation function 365 calculates the reliability using the total angle range.

Alternatively, the calculation function 365 is the number and photon counting when it is assumed that the number of X-ray photons counted by the photon counting function 161 is proportional to the number of X-ray photons incident on the detection element at each point of the subject P. The reliability is calculated by dividing an angle range in which the difference from the number of X-ray photons counted by the function 161 exceeds a predetermined threshold value by the entire angle range.

The calculation function 365 is a difference between the number of X-ray photons shown by the dotted line D in FIG. 6 and the number of X-ray photons shown by the solid line S in FIG. 6 at each point of the subject P. Specifies an angle range where is above a predetermined threshold. This predetermined threshold value determines whether or not the number of X-ray photons corrected by the correction function 364 is reliable. This predetermined threshold value is determined, for example, by the calculation function 365. If this difference is less than a predetermined threshold, the number of X-ray photons corrected by the correction function 364 is determined to be reliable. When this difference is equal to or greater than a predetermined threshold value, the number of X-ray photons corrected by the correction function 364 is determined to be unreliable. Then, the calculation function 365 calculates the above-mentioned reliability.

Alternatively, the calculation function 365 is the number and photon counting when it is assumed that the number of X-ray photons counted by the photon counting function 161 is proportional to the number of X-ray photons incident on the detection element at each point of the subject P. The reliability may be calculated by dividing an angle range in which the difference from the number of X-ray photons counted by the function 161 is less than a predetermined threshold value by an angle range in which the difference is equal to or more than a predetermined threshold value.

Alternatively, the calculation function 365 is the number and photon counting when it is assumed that the number of X-ray photons counted by the photon counting function 161 is proportional to the number of X-ray photons incident on the detection element at each point of the subject P. The reliability is the value obtained by dividing the value obtained by integrating the difference in the angle range in which the difference from the number of X-ray photons counted by the function 161 is less than a predetermined threshold value by the value obtained by integrating the difference in the entire angle range. calculate.

First, the calculation function 365 determines the number of X-ray photons shown by the dotted line D in FIG. 6 and the number of X-ray photons shown by the solid line S in FIG. 6 at each point of the subject P. Specify an angular range in which the difference between the two exceeds a predetermined threshold. Then, the calculation function 365 calculates the above-mentioned reliability.

Alternatively, the calculation function 365 is the number and photon counting when it is assumed that the number of X-ray photons counted by the photon counting function 161 is proportional to the number of X-ray photons incident on the detection element at each point of the subject P. The value obtained by integrating the difference in the angle range in which the difference from the number of X-ray photons counted by the function 161 is less than the predetermined threshold value is divided by the value obtained by integrating the difference in the angle range in which the difference is equal to or more than the predetermined threshold value. The value may be calculated as the reliability.

Even when the calculation function 365 calculates the above-mentioned reliability instead of the reliability calculated in step S9, the photon counting type X-ray CT apparatus 1 has the above-mentioned effect.

The above-mentioned processors include, for example, a CPU (Central Processing Unit), a GPU (Graphics Processing Unit), an application specific integrated circuit (ASIC), a programmable logic device (PLD), and a field programmable gate array. (Field Programmable Gate Array: FPGA). Further, the programmable logic device (Programmable Logic Device: PLD) is, for example, a simple programmable logic device (Simple Programmable Logic Device: SPLD) or a compound programmable logic device (Complex Programmable Logic Device: CPLD).

In the above-described embodiment, the high voltage generation circuit 11, the collimator adjustment circuit 12, the gantry drive circuit 13, the data acquisition circuit 16, the sleeper drive circuit 22, and the processing circuit 36 read out and execute the program stored in the storage circuit 35. By doing so, the function was realized, but the function is not limited to this. Instead of storing the program in the storage circuit 35, the program may be directly incorporated in each of these circuits. In this case, these circuits realize their functions by directly reading and executing the embedded program.

Each circuit shown in FIG. 1 may be appropriately distributed or integrated. For example, the processing circuit 36 is a scan control circuit that executes each of the scan control function 361, the preprocessing function 362, the image generation function 363, the correction function 364, the calculation function 365, the display control function 366, and the control function 367, and the preprocessing. It may be distributed to a circuit, an image generation circuit, a correction circuit, a calculation circuit, a display control circuit, and a control circuit. Further, the data acquisition circuit 16 is distributed to a photon counting circuit, a dosimetry circuit, a projection data generation circuit and a conversion circuit that execute the respective functions of the photon counting function 161 and the dose measuring function 162, the projection data generation function 163 and the conversion function 164. May be done. Further, the high voltage generation circuit 11, the collimator adjustment circuit 12, the gantry drive circuit 13, the data acquisition circuit 16, the sleeper drive circuit 22, and the processing circuit 36 may be arbitrarily integrated.

Further, the processing other than the above-mentioned step S1 may be performed by the image processing apparatus instead of the photon counting type X-ray CT apparatus 1. This image processing device has a correction function and a calculation function. The correction function is for the energy band set on the energy distribution of the X-rays emitted by the X-ray tube 141, the position of the X-ray tube 141, and the X-rays counted for each detection element that detects the incident X-ray photons. The number of photons is corrected based on the detection characteristics of the detection element. The calculation function calculates the reliability of the pixels of the reconstructed image based on the result corrected by the correction function.

According to at least one embodiment described above, the diagnosis using the reconstructed image can be appropriately performed.

Although some embodiments of the present invention have been described, these embodiments are presented as examples and are not intended to limit the scope of the invention. These embodiments can be implemented in various other embodiments, and various omissions, replacements, and changes can be made without departing from the gist of the invention. These embodiments and variations thereof are included in the scope of the invention described in the claims and the equivalent scope thereof, as are included in the scope and gist of the invention.

141 X-ray tube 15 Detector 161 Photon counting function 364 Correction function 365 Calculation function 32 Display

Claims (9)

An X-ray tube that irradiates the subject with X-rays,
A detector having a plurality of detection elements for detecting incident X-ray photons, and a detector.
An energy band set on the X-ray energy distribution, a photon counting unit that counts the number of X-ray photons for each X-ray tube position and detection element, and a photon counting unit.
A correction unit that corrects the number of X-ray photons counted by the photon counting unit based on the dose of the X-ray incident on the detection element and the detection characteristics of the detection element.
The reciprocal of the error generated when the correction unit corrects is compared with the threshold value, and the pixels of the reconstructed image are based on the range of the incident angle of the X-ray dose when the reciprocal of the error exceeds the threshold value. A calculation unit that calculates the reliability and
A photon counting type X-ray CT apparatus. The correction unit converts the number of X-ray photons counted by the photon counting unit into the number of X-ray photons incident on the detection element by the number of X-ray photons counted by the photon counting unit. The photon counting type X-ray CT apparatus according to claim 1, which is corrected to a number when it is assumed to be proportional. The photon counting type X-ray CT apparatus according to claim 1 or 2, wherein the calculation unit further calculates a value used for diagnosis and associates the value with the reliability. The photon counting type X-ray CT apparatus according to any one of claims 1 to 3, further comprising a display unit for displaying the reliability. The photon counting type X-ray CT apparatus according to claim 4, wherein the display unit displays the reliability together with the reconstructed image. The photon counting type X-ray CT apparatus according to claim 5, wherein the display unit displays the reliability on the reconstructed image. The photon counting type X-ray CT apparatus according to claim 5, wherein the display unit displays the reliability when the displayed cursor moves. When the statistic that is an index of the magnitude of the reliability of the pixel included in the region of interest of the reconstructed image is smaller than a predetermined threshold value, the X-ray tube has a lower dose than the X-ray. The photon counting type X-ray CT apparatus according to any one of claims 1 to 7, further comprising a scan control unit for irradiating the subject and scanning the subject. The X-ray photons counted for each detection element that detects the energy band set on the X-ray energy distribution irradiated by the X-ray tube, the position of the X-ray tube, and the incident X-ray photons. A correction unit that corrects the number of X-rays incident on the detection element based on the dose of the X-ray and the detection characteristics of the detection element.
The reciprocal of the error generated when the correction unit corrects is compared with the threshold value, and the pixels of the reconstructed image are based on the range of the incident angle of the X-ray dose when the reciprocal of the error exceeds the threshold value. A calculation unit that calculates the reliability and
An image processing device.

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