Imaging Device, Method, and Program

Abstract

To optimize an energy bin.SOLUTION: An imaging apparatus according to an embodiment comprises a processing circuit. The processing circuit estimates, for two or more detector elements in a photon-counting detector, respective energy spectra of an X-ray beam incident on the corresponding detector elements from an X-ray source, the energy spectra being estimated by modeling X-ray attenuation as a function of X-ray energy. For each detector element of the two or more detector elements, the processing circuit sets a first energy threshold corresponding to the maximum value of a first energy range that is detected by the detector element on the basis of the energy spectra.SELECTED DRAWING: Figure 2

Description

Technical Field

[0001] Embodiments of the present invention relate to an imaging apparatus, method, and program.

Background Art

[0002] CT (Computed Tomography) systems and methods are widely used, particularly for medical imaging and medical diagnosis. A CT system generally creates projection images of one or more cross-sectional slices of a subject's body. A radiation source such as an X-ray source (sometimes simply referred to as a source) irradiates the body with radiation from one side of the body. Since a collimator generally adjacent to the X-ray source restricts the angular range of the X-ray beam, the radiation impinging on the body is substantially restricted to a planar region (i.e., the X-ray projection plane) that defines the cross-sectional slice of the body. At least one detector (generally a number of detectors) on the opposite side of the body receives the radiation that has passed through the body in the projection plane. By receiving the electrical signal output from the detector and processing this electrical signal, the attenuation of the radiation that has passed through the body is measured. A multi-slice detector that performs volume projection of the body instead of planar projection may be used.

[0003] Generally, an X-ray source is mounted on a gantry that rotates around the long axis of the body. The detector is similarly mounted on the side of the gantry opposite the X-ray source. Projection attenuation measurements are performed at a series of gantry rotation angles, and the projection data / sinogram is sent to a processing device (processor) via a slip ring provided between the gantry rotor and stator, and then the projection data is processed using a CT reconstruction algorithm (e.g., inverse Radon transform, filtered back projection, cone beam reconstruction based on Feldkamp, iterative reconstruction, or other methods) to obtain a cross-sectional image of the body. For example, the reconstructed image can be a digital CT image that is a square matrix of elements (pixels). In this case, each of these elements represents a volume element (volume pixel or voxel) of the patient's body. Depending on the CT system, the X-ray source passes through a spiral or helical trajectory relative to the body by a combination of the translation of the body and the rotation of the gantry relative to the body. Then, multiple views are used to reconstruct a CT image showing the internal structure of one slice or multiple slices.

[0004] An Energy-Integrating Detector (EID) has been used to measure CT projection data. As an alternative method, a Photon-Counting Detector (PCD) has a number of advantages, including the ability to perform spectral CT and the function of enhancing resolution by dividing the scan area into a number of smaller "pixels" of the detector, compared to the energy-integrating detector. Semiconductor-based PCDs bring unique benefits to spectral CT but also pose inherent challenges. For example, a PCD can decompose the detected X-rays into each of a plurality of energy bins corresponding to each of a plurality of energy ranges. The energy range of each bin is the same for all elements (detection elements) within the detector. However, the optimal energy range for the detection elements in a part of the X-ray beam is not necessarily optimal for the detection elements in another part of the X-ray beam. This is because different parts of the X-ray beam can have different X-ray energy spectra. For example, a plurality of different parts of the X-ray beam may have a plurality of different energy spectra from each other. Therefore, using the same range of energy bins for all detection elements may degrade performance and image quality. Thus, an excellent method for optimizing these ranges (e.g., energy thresholds and the size of the energy bins) as a function of the position within the X-ray beam is desired.

Prior Art Documents

Patent Documents

[0005]

Patent Document 1

Summary of the Invention

Problems to be Solved by the Invention

[0006] The problem to be solved by the present invention is to optimize the energy bins.

Means for Solving the Problems

[0007] The imaging device according to the embodiment includes a processing circuit. The processing circuit estimates, for two or more detection elements of a photon counting detector, the respective energy spectra of the X-ray beams of an X-ray source incident on the corresponding detection elements by modeling X-ray attenuation as a function of X-ray energy. The processing circuit sets, for each detection element of the two or more detection elements, a first energy threshold corresponding to the maximum value of a first energy range detected by each detection element based on the energy spectrum.

Brief Description of the Drawings

[0008]

Figure 1A

Figure 1B

Figure 1C

Figure 1D

Figure 2

Figure 3

Figure 4

DETAILED DESCRIPTION OF THE INVENTION

[0009] The description set forth below in connection with the accompanying drawings is intended to describe various aspects of the disclosed subject matter and is not necessarily intended to show only those aspects. In some cases, the description includes specific details for the purpose of providing an understanding of the disclosed subject matter. However, it will be apparent to those skilled in the art that the various aspects may be practiced without these specific details. In some cases, well-known structures and components are shown in block diagram form to avoid obscuring the concepts of the disclosed subject matter.

[0010] References throughout the specification to "one aspect" or "an aspect" mean that a particular feature, structure, characteristic, operation, or function described in connection with that aspect is included in at least one aspect of the disclosed subject matter. Thus, the phrases "in one aspect" or "in an aspect" in the specification do not necessarily refer to the same aspect. Further, the particular features, structures, characteristics, operations, or functions may be combined in any suitable manner in one or more aspects. Additionally, aspects of the disclosed subject matter are intended to include and encompass modifications and variations of the described aspects.

[0011] It should be noted that, as used in this specification and the claims, unless the context otherwise indicates, the singular forms may include plural referents. That is, unless specifically specified otherwise, as used in this application, the terms "a", "an" and similar terms have the meaning of "one or more". Also, of course, terms such as "top", "bottom", "front", "rear", "side", "inner", "outer", etc., when used in this application, only indicate a reference point and do not necessarily limit the features of the disclosed subject matter to a specific orientation or arrangement. Furthermore, terms such as "first", "second", "third", etc., are only used to identify one of several parts, components, reference points, operations and / or functions described in this application, and similarly, do not necessarily limit the features of the disclosed subject matter to a specific arrangement or orientation.

[0012] The embodiments described in this application relate to, for example, the setting of a changeable energy threshold of a detection element (pixel) in a photon counting detector. For example, for each of a plurality of detection elements of a photon counting detector, it relates to a computed tomography (CT) system, method and program for optimally adjusting a plurality of energy thresholds according to the energy spectrum at corresponding points in an X-ray beam. In the following description, the photon counting detector may sometimes be simply referred to as a detector.

[0013] In photon counting CT, the PCD can detect the energy of each incident photon by direct conversion, for example, by a CdTe or CZT energy-resolving photon counting spectral detector. Compared with CT using an EID, CT using a PCD provides additional energy-dependent information in its measurements. This information enables material discrimination and spectral imaging.

[0014] Front-end electronic devices of a PCCT system, such as an Application Specific Integrated Circuit (ASIC) for a specific purpose, convert an analog signal, compare it with a predetermined threshold (voltage), and thereby obtain a measured value of the count of a large number of energy bins. The photon counting ASIC design may include, for example, 2 to 6 energy thresholds with 2 to 6 energy bins. In some examples, a static value (e.g., a fixed value) within the range of 25 to 50 keV is used as the first threshold by some systems. However, since the threshold depends on the system-specific configuration settings, there is no established theory on how to determine the optimal energy threshold. In particular, the lowest energy bin may contain mixed data including attenuated primary X-ray beams and scattered X-rays. On the other hand, the photoelectric interaction is the main attenuation process for X-rays in the low energy range of the X-rays, and thereby low energy X-rays show higher sensitivity to differences in substances. On the other hand, the X-ray scattering X-ray process results in scattered X-rays with lower energy than the incident X-rays. As a result, most of the scattered X-rays are found in the low energy range, which, unless X-ray scattering correction is performed, affects the measurement adversely and causes bias / artifacts in the CT image. The selection of the optimal low energy cut-off threshold is in a trade-off relationship between removing unnecessary noise due to scattered X-rays by increasing the cut-off threshold and collecting as much signal as possible at low X-ray energies by reducing the cut-off threshold. If the optimal energy threshold is not determined and the appropriate energy bin settings are not made, valuable data (useful data) will be excluded from the measurement. Therefore, the method described in the present application optimizes the performance of the PCCT system and enhances its clinical effect by introducing dynamic and position-dependent determination of the energy threshold and bin settings into the PCCT system.

[0015] Reference will now be made to a plurality of drawings, in which like reference numerals refer to the same or corresponding parts throughout several views. FIG. 1A shows a cross-sectional view of an object (e.g., a subject, a specimen, a patient, a torso, a phantom, etc.) and the trajectories of X-rays emitted from a radiation source and passing through the object en route to respective pixels or elements in a detector. For simplicity, only half of the X-ray trajectories are shown. The other half can be obtained by folding the shown trajectories about a center line.

[0016] Some of the X-ray trajectories pass through only soft tissue, while other X-ray trajectories pass through bone (e.g., ribs) in addition to passing through soft tissue. Further, for some source positions, such as directly above the patient, some of the X-ray trajectories pass through only the patient's arm, while other trajectories pass through the torso. As a natural result, for other source positions, such as on the left or right side of the patient, some of the X-ray trajectories travel in a lateral direction (sideways direction) and pass first through the first arm, then the torso, and subsequently through the second arm. Thus, the distance traveled by the X-rays and the material through which the X-rays pass vary depending on the various positions of the source. In the case of a completely circular cylinder, some of the X-ray trajectories pass through only a small amount of tissue, for example, at the periphery, while other X-ray trajectories pass through the center of the object or patient. As can be seen from these examples, there is a difference in attenuation between the central and peripheral X-rays.

[0017] Furthermore, in order to limit the radiation dose by shaping or narrowing the X-ray beam, a dose compensation filter (bowtie filter) may be inserted between the radiation source and the patient. FIG. 1B shows, in one aspect of the present disclosure, a cross-sectional view of the torso and the trajectory of the fan-shaped X-ray when a dose compensation filter is provided. In one aspect, the addition of the dose compensation filter reduces the amount of X-ray directed towards the periphery of the patient. Therefore, the addition of the dose compensation filter can reduce the radiation dose to the patient. For example, when performing a head CT scan, a smaller fan angle may be desirable in order to keep the X-ray dose as low as reasonably acceptable. Similarly, in the case of a cardiac CT scan, it may be desirable to reduce the radiation dose to the patient by reducing the radiation dose to clinically irrelevant body parts such as the arms. This can be achieved by a dose compensation filter having a shape where the edge of the dose compensation filter is thicker than the central portion. In this case, the dose compensation filter is arranged such that the X-ray directed towards the periphery of the fan angle passes through the thicker edge of the dose compensation filter, and the X-ray (central X-ray) directed towards the center of the fan angle passes through the thinner center of the dose compensation filter. Filtering of the beam (pre-filtering) by the dose compensation filter results in different spectra along the fan angle. Therefore, in the absence of the scanned object or the patient, the peripheral detection elements, which are the peripheral portions of the photon counting detector, receive a spectrum that is shifted to a higher energy compared to the detection elements closer to the center.

[0018] For various clinical applications and various patient sizes, various beam shapes / widths can be beneficial. Therefore, an appropriate dose compensation filter can be selected from a discrete set of dose compensation filters based on the specific application and the patient. In one embodiment, the energy spectrum as a function of the fan angle can be calibrated (corrected) and stored for each appropriate dose compensation filter in the discrete set of dose compensation filters.

[0019] FIG. 1C shows an example of a beam spectrum after passing through a dose compensation filter at various fan angles. In FIG. 1C, the edge X-ray 105 (see FIG. 1B) generates an edge spectrum 105a, and the central X-ray 115 (see FIG. 1B) generates a central spectrum 115a. Further, the intermediate X-ray 110 (see FIG. 1B) generates an intermediate spectrum 110a. Due to the energy-dependent attenuation of the dose compensation filter, the edge spectrum 105a is shifted in the direction of higher energy than the intermediate spectrum 110a, and the intermediate spectrum 110a is shifted in the direction of higher energy than the central spectrum 115a.

[0020] FIG. 1D shows an example of an energy bin 120 applied to a beam spectrum after passing through a dose compensation filter at various fan angles. For example, the threshold value defining the energy bin 120 is set to a predetermined energy level. For example, the threshold value is set every 20 keV to obtain a first energy bin 120a, a second energy bin 120b, a third energy bin 120c, a fourth energy bin 120d, and a fifth energy bin 120e. It will be understood by those skilled in the art that various threshold levels and energy bin sizes can be used.

[0021] It should be noted that the pixels detecting the edge X-ray 105 are substantially not exposed to X-rays with energies lower than 45 keV, while the pixels detecting the central X-ray 115 are exposed to X-rays with energies down to a low range of 20 keV. Therefore, as shown in FIG. 1D, the first energy bin 120a does not substantially provide spectral information regarding the edge X-ray 105 because the edge spectrum 105a disappears at 45 keV.

[0022] Therefore, the X-rays with energies lower than 45 keV incident on the pixels of the detector corresponding to the edge X-ray 105 are likely to be from scatter rather than from the primary X-ray beam. As a result, for the pixels of this detector, the low energy cut-off threshold of the first energy bin 120a can be, for example, 50 keV.

[0023] The first energy bin mentioned here refers to, for example, the energy bin with the lowest energy among a plurality of energy bins set for a certain pixel. The low energy cut-off threshold refers to, for example, the lowest energy within the energy range of the first energy bin. The cut-off threshold is also referred to as the first energy threshold.

[0024] In some embodiments, this low energy cut-off threshold can be slightly lowered (for example, to 45 keV) considering the finite energy resolution of the PCD (for example, this energy resolution is 5 - 10 keV). Also, in some embodiments, this low energy cut-off threshold can be slightly increased considering the attenuation (additional attenuation) resulting from the propagation within the patient's body, especially in the case of a large and / or high-density patient. Note that the additional attenuation refers to, for example, the attenuation that occurs further when an object such as a patient is present compared to the case when no object is present.

[0025] Regarding the pixel of the detector corresponding to the intermediate spectrum 110a, the low energy cut-off threshold can be, for example, 40 keV. Similar to the low energy cut-off threshold of the pixel corresponding to the edge spectrum 105a, this 40 keV cut-off threshold can be adjusted upward or downward according to the finite energy resolution of the PCD and the additional attenuation by the patient.

[0026] Regarding the pixel of the detector corresponding to the central spectrum 115a, the low energy cut-off threshold can be, for example, 25 keV. Similar to the low energy cut-off threshold of the pixel corresponding to the edge spectrum 105a, this 25 keV cut-off threshold can be adjusted upward or downward according to the finite energy resolution of the PCD and the additional attenuation by the patient.

[0027] In view of the above, it is possible to perform simulations (or measurements / calibrations in some embodiments) of the X-ray spectrum at each position along the pixel array of the PCD. Then, these X-ray spectra can be used to determine the low-energy cut-off thresholds at each position along the pixel array of the PCD. That is, in order to optimize the setting of the threshold, a variable energy threshold setting can be applied to various pixels. That is, various pixels can be configured so that a variable energy threshold can be set. Each individual pixel or group of pixels (e.g., a PCD module consisting of 100 pixels) has an energy threshold automatically set by a settable first energy threshold, an additionally settable energy threshold, and a method of automatically setting the energy threshold based on various inputs.

[0028] The additionally settable energy threshold refers to, for example, an energy threshold other than the first energy threshold and is used to define the boundary of the energy bin. For example, in the example of FIG. 1D, the additionally settable energy thresholds are the energy thresholds that define the boundaries of the respective energy bins of the plurality of energy bins 120a to 120e. More specifically, the plurality of additionally settable energy thresholds include the energy threshold (second energy threshold) between the first energy bin 120a and the second energy bin 120b, the energy threshold (third energy threshold) between the second energy bin 120b and the third energy bin 120c, the energy threshold (fourth energy threshold) between the third energy bin 120c and the fourth energy bin 120d, the energy threshold (fifth energy threshold) between the fourth energy bin 120d and the fifth energy bin 120e, and the maximum value of the energy range of the fifth energy bin 120e (sixth energy threshold). The plurality of additionally settable energy thresholds are simply also referred to as additional energy thresholds.

[0029] For example, when the measured energy of the X-rays detected by one detection element is less than the second energy threshold and greater than the first energy threshold, it is discriminated that the detected X-rays are included within the first energy bin 120a. The same applies to other energy bins.

[0030] FIG. 2 is a flowchart showing a method 200 for setting energy thresholds by beam simulation and calibration according to one aspect of the present disclosure. In one aspect, the first energy threshold set for a plurality of pixels can be set by using a beam spectrum estimated based on energy-dependent attenuation caused by passing through a dose compensation filter. In step S201, a simulation of the spectrum of the beam emitted from the radiation source (X-ray tube spectrum) is performed, and the X-ray tube spectrum obtained by the simulation is combined with the attenuation of the beam spectrum (estimated attenuation) that has passed through a predetermined filter. The attenuation can be estimated by calculating based on filter characteristics such as shape and material. For example, the predetermined filter can be a large dose compensation filter for adult patients, a small dose compensation filter for pediatric patients, or other filters installed in a CT system for specific clinical applications.

[0031] According to embodiments, the X-ray tube spectrum of the beam emitted from the radiation source may be based on a calibration scan. As another method, the X-ray tube spectrum of the beam emitted from the radiation source may be based on a simulation. For example, the X-ray tube spectrum can be obtained by performing a simulation for obtaining the X-ray tube spectrum based on the setting information of the radiation source (information including information such as tube voltage and tube current, for example, the X-ray tube 501 (see FIG. 4) described later).

[0032] That is, in step S201, a simulation of the X-ray tube spectrum is performed, and the beam spectrum is estimated (calculated) using the X-ray tube spectrum obtained by the simulation.

[0033] Thus, in step S201, for example, for two or more detection elements of the photon counting detector, the respective energy spectra of the X-ray beams of the X-ray source incident on the corresponding detection elements are estimated by modeling the X-ray attenuation as a function of the X-ray energy. Specifically, the respective energy spectra are estimated by modeling the attenuation of the X-rays when the X-ray beam passes through the filter. Thus, for example, the respective energy spectra are estimated using a model of the attenuation of the X-rays when the X-ray beam passes through the filter (the model obtained by the above modeling).

[0034] In step S203, optionally, the pixels in the detector can also be grouped according to a range desired by the user of the beam fan angle. For example, taking a detector of 896 rows / columns, 32 groups consisting of 28 rows / columns covering the entire beam fan angle can be formed. As another method, the pixels can also be grouped by module. Furthermore, each pixel can be its own group. That is, one group may be composed of one pixel. Note that the pixels may or may not be grouped. For example, when grouped, the process of step S205 described below is performed for each group of pixels. On the other hand, when not grouped, the process of step S205 is performed for each individual pixel (one pixel).

[0035] In step S205, for each group of pixels or for each pixel, a first energy threshold (i.e., a low energy cut-off threshold) can be determined based on the calculated beam spectrum. Since the energy spectrum is unique for each group or each pixel, in this way the first energy threshold can be made group-specific or pixel-specific.

[0036] For example, in step S205, for each detection element of two or more detection elements, a first energy threshold corresponding to the maximum value of the first energy range detected by each detection element is set based on the energy spectrum. The first energy range referred to here is the range of energy that is not counted (not counted) from 0 keV shown in FIG. 1D or the like to the first energy threshold.

[0037] Also, in steps S203 and S205, the detection elements are grouped into a plurality of groups, where each group of the plurality of groups has an energy spectrum of each X-ray beam.

[0038] In some embodiments, the central spectrum 115a has a minimum energy Emin of 20 keV. Accordingly, the first energy threshold of the pixel (or group of pixels) that detects the central X-ray 115 is set to 20 keV. In another aspect, the PCD has an energy resolution a. And the first energy threshold can be adjusted by an amount based on the energy resolution a. For example, the first energy threshold can be determined by subtracting the energy resolution from the minimum energy (e.g., Emin - a). That is, when the energy resolution is 5 keV and the minimum energy of the central spectrum 115a is 20 keV, the first energy threshold is set to 15 keV.

[0039] The minimum energy Emin can be determined using various methods. In some embodiments, the minimum energy can be the energy for a measured spectrum that exceeds a predetermined threshold. For example, this predetermined threshold can be a value based on an absolute scale or a relative value. This predetermined threshold can be related to the peak of the energy spectrum or the integrated value of the energy spectrum. For example, this predetermined threshold can be the peak of the energy spectrum or the integrated value of the energy spectrum. In this case, the energy spectrum is represented by a count rate, irradiance, intensity, radiation flux (rate), etc. As another method, the minimum energy can be based on the integration of the energy spectrum. For example, the minimum energy can be selected to remove a predetermined percentage (e.g., 2% or 5%) of the energy spectrum, which is measured, for example, by counts or energy. In some embodiments, the minimum energy is determined using both the energy spectrum and an estimated value of a noise signal such as a scattered x-ray flux.

[0040] In another aspect, other energy thresholds (e.g., the threshold between the first energy bin 120a and the second energy bin 120b (the second energy threshold) or the threshold between the second energy bin 120b and the third energy bin 120c (the third energy threshold), etc.) can be determined / adjusted based on the position of the pixel.

[0041] To determine the respective ranges of the energy bins, various rules can be used. In one example shown in FIG. 1D, each of the plurality of energy bins 120 spans the same energy range (e.g., 20 keV). As another method, the width of the energy bin (spectral width) can be selected to achieve a noise balance. Further, the width of the energy bin can also be selected so as to equalize the count number or the signal-to-noise ratio in each of the plurality of energy bins. Also, the width of the energy bin can be selected based on a first energy threshold or a reference table (look-up table (LUT)) according to some predetermined formula. Further, other methods can be used to select the width of the energy bin. The width of the energy bin may not be uniform.

[0042] For example, the first energy bin 120a of the central spectrum 115a may have a width of 40 keV and span from 20 keV to 60 keV. The second energy bin 120b may have a width of 20 keV and span from 60 keV to 80 keV. The third energy bin 120c may have a width of 10 keV and span from 80 keV to 90 keV. The fourth energy bin 120d may have a width of 10 keV and span from 90 keV to 100 keV. The fifth energy bin 120e may have a width of 20 keV and span from 100 keV to 120 keV. This enables the first energy bin 120a to measure all low-energy X-rays including those below 40 keV. Similarly, the adjusted third energy bin 120c and fourth energy bin 120d more accurately resolve X-rays in the range of 80 - 100 keV.

[0043] For all groups of pixels or all pixels, various first energy thresholds and additional energy thresholds are calculated for each dose compensation filter (for example, the large dose compensation filter described above, the small dose compensation filter described above, etc.), and the results (the first energy threshold and the additional energy threshold) may be stored in a reference table. For example, in the reference table, an identifier for identifying a group of pixels or a pixel, an identifier for identifying a dose compensation filter, and the first energy threshold and the additional energy threshold may be registered in association with each other. The reference table is stored as a scan setting (scan condition) set in advance in this way and is called based on the filter used for a given scan (for example, an imaging scan). For example, values may be stored as a pre-created table in various formats such as a reference table (LUT). Before performing a scan (for example, an imaging scan), the setting of the corresponding energy threshold may be applied to the pixels based on the predetermined filter used. In particular, this method may be executed during the calibration phase for a new filter before guiding the patient onto the gantry. As a result, the waiting time (patient time) during the scan is shortened.

[0044] FIG. 3 shows a flowchart of an example of a method 300 for setting an energy threshold by beam simulation and pre-scanning. In some embodiments, a scout scan is performed to estimate the shortest path length (length of the shortest path) between the radiation source and the patient at various predetermined orientations (directions) around the patient. In this method, the setting of the first energy threshold of the pixels is not based only on the design of a predetermined filter (e.g., filter characteristics). The first energy threshold is further optimized by estimating the shape of the patient (such as the shape of the cross-section of the patient) during the execution of a scout scan performed prior to the imaging scan (this scan). Next, the first energy threshold is adjusted based on the shortest path lengths determined for each view angle (i.e., the orientation / position of the x-ray source). That is, simulations can be performed to obtain an energy spectrum based on the combination of the spectrum calculated based on the attenuation resulting from the shortest path length and beam filtering at each view angle.

[0045] For example, the cross-section of the patient's torso may be asymmetric and oval-shaped. In this case, the transverse diameter (e.g., the axis extending horizontally with respect to the ground) is longer than the longitudinal diameter. The longitudinal axis is, for example, orthogonal to the transverse axis. That is, the width of the torso from arm to arm is longer than the thickness of the torso from the spine to the front of the chest. Therefore, the shortest path length along the transverse axis from the radiation source to the outer surface of the patient is shorter than the shortest path length along the longitudinal axis from the radiation source to the outer surface of the patient.

[0046] In some embodiments, similar to method 200, in step S301, a simulation of the x-ray tube spectrum is performed, and then the beam spectrum passing through the predetermined filter is calculated based on the predetermined filter characteristics.

[0047] In step S303, as an option, pixels in the detector are grouped based on the calculated beam spectrum. Note that, similar to the process of step S203 described above, in step S303, the pixels may or may not be grouped. For example, if grouped, for each group of pixels, the processes of steps S309 and S311 described below are performed. On the other hand, if not grouped, for each individual pixel (one pixel), the processes of steps S309 and S311 are performed.

[0048] In step S305, the patient is irradiated with radiation at a low dose in a predetermined orientation around the patient by scout scanning. To minimize the radiation dose, the scout scan is performed at a narrower view angle than the imaging scan. For example, by scout scanning, the patient is irradiated with radiation at a low dose along the vertical axis and, separately, the patient is irradiated with radiation at a low dose along the horizontal axis. To estimate the cross-sectional shape of the patient, the scout scan irradiates the patient with radiation by any number of low-dose scans in various orientations.

[0049] In step S307, the shortest path length is estimated based on the scout scan. In step S309, the first energy threshold is determined using the shortest path length determined from the scout scan together with the beam spectrum after passing through a predetermined filter (the calculated beam spectrum).

[0050] For example, the scout scan finds that the shortest path length along the horizontal axis is shorter. Therefore, the beam is further attenuated as it passes through more substances (for example, compared to the case of a circular object). Thus, the calibration adjusts the first energy threshold to a higher energy for all groups of pixels or all individual pixels based on the shortest path length determined by the scan performed on the horizontal axis.

[0051] In another aspect, using the calculated post-filter spectrum (i.e., the spectrum after passing through a predetermined filter) along with the shortest path length measured by a scout scan, a first energy threshold is determined at various view angles (i.e., various orientations of the X-ray source around the patient). That is, the first energy threshold is determined at each view angle based on the shortest path length in the corresponding orientation. For example, in the case of a source oriented on the horizontal axis (a source oriented in the direction of the horizontal axis), it is found by a scout scan that the shortest path length occurs in that orientation. Thus, the first energy threshold of the central X-ray 115 for a source oriented on the horizontal axis is set to 20 keV. Subsequently, the source is rotated by a predetermined angle, for example, 1 degree clockwise. By this scout scan, it is found that the shortest path length at a view angle 1 degree greater than the horizontal axis (assuming the view angle when the source is on the horizontal axis is 0 degrees, a view angle of 1 degree) is slightly longer than the scout scan on the horizontal axis. Based on such measurements, the first energy threshold of the central X-ray 115 from a source oriented 1 degree above the horizontal axis orientation is set to, for example, 21 keV or, preferably, a value higher than 20 keV.

[0052] In some embodiments, when the view angle of the imaging scan is different from the view angle of the scout scan, for the shortest path length at each view angle of the imaging scan, the CT system can estimate the shortest path length by interpolation (or extrapolation using the shortest path length of the nearest scout scan view angle and the shortest path length of the second nearest scout scan view angle) between the shortest path length of the nearest scout scan view angle and the shortest path length of the second nearest scout scan view angle. Further, when the measured / simulated spectra at both ends are available, an energy spectrum corresponding to the view angle and the position along the fan angle (e.g., the position along the direction in which the detector detection elements are arranged) can be determined by interpolation or extrapolation.

[0053] In another aspect, during the execution of a scout scan prior to an imaging scan, the shape of the patient is estimated, and the additional energy threshold can be further optimized by adjusting the additional energy threshold based on the shortest path length at each view angle and the beam spectrum calculated based on beam filtering. That is, based on the orientation of the radiation source, the size of the energy bin 120 of a group of pixels or individual pixels at that specific orientation can be adjusted. In this way, for example, from the data obtained by the scout scan, an additional energy threshold can be set in advance for each view angle.

[0054] In another aspect, the respective first energy threshold and additional energy threshold of a group of pixels or individual pixels can be adjusted according to the view angle. For example, during a scan when the radiation source is oriented along the horizontal axis, the intermediate X-ray 110 does not pass through the patient. During a scan when the radiation source is oriented along the vertical axis, since the cross-section of the patient is in an oval shape, the intermediate X-ray 110 passes through the patient. Thus, the orientation of the radiation source around the patient affects the detected spectrum of the beam and the possible attenuation detected by the pixels that detect the intermediate X-ray 110 and its position along the fan angle. In this way, for example, since the path length changes depending on the shape of the patient, the first energy threshold and the additional energy threshold can be adjusted (set) for each combination of the view angle and the fan angle.

[0055] Some settings of the first energy threshold and settings of the additional energy threshold are predefined. And thus, calibration of the predefined settings of the first energy threshold and the additional energy threshold can be generated. For example, information (calibration information) for calibrating the predefined settings of the first energy threshold and the additional energy threshold can be generated. This calibration (calibration information) can be stored in, for example, a LUT.

[0056] Thus, in step S309, the energy spectrum is estimated for each group of pixels or each pixel by modeling the attenuation of X-rays when the X-ray beam passes through the filter and the object. Thus, for example, the energy spectrum is estimated using a model of the attenuation of X-rays when the X-ray beam passes through the filter and the object (the model obtained by the above modeling). Then, in step S309, a first energy threshold is determined based on the estimated energy spectrum.

[0057] Also, in step S309, the respective energy spectra are estimated by modeling the energy spectrum based on the simulation of the X-ray beam passing through the filter and the object. Thus, for example, the respective energy spectra are estimated using a model of the energy spectrum based on the simulation of the X-ray beam passing through the filter and the object (the model obtained by the above modeling).

[0058] Also, in step S309, the respective energy spectra are estimated by modeling the energy spectrum based on the calibration of the X-ray beam passing through the filter and the simulation of the X-ray beam passing through the object. Thus, for example, the respective energy spectra are estimated using a model of the energy spectrum based on the calibration of the X-ray beam passing through the filter and the simulation of the X-ray beam passing through the object (the model obtained by the above modeling).

[0059] Also, in step S309, the energy spectrum of the X-ray beam is estimated by determining the shortest path length, which is the length of the shortest path from the source of the X-ray beam passing through the object to the object at one view angle. Then, the shortest path length is used to perform a simulation of the X-ray beam passing through the object at one view angle.

[0060] Also, in step S309, the energy spectrum of the X-ray beam is estimated by determining the shortest path length, which is the length of the shortest path from the source of the X-ray beam passing through the object to the object at each view angle. Then, using the shortest path length at each view angle, a simulation of the X-ray beam passing through the object at the corresponding view angle is performed.

[0061] Also, in step S309, the energy spectrum is estimated by modeling the X-ray attenuation after the X-ray passes through the object using a scout scan that acquires scout projection data using a dose lower than the dose used in subsequent imaging scans. Thus, for example, the energy spectrum is estimated using a model of X-ray attenuation after the X-ray passes through the object (the model obtained by the above modeling).

[0062] In step S311, the CT system determines the closest first energy threshold from the LUT for each group or each pixel of the pixels in the detector and applies the first energy threshold to the imaging scan and subsequent data processing. Further, the CT system determines the closest additional energy threshold from the LUT for each group or each pixel of the pixels in the detector and applies the additional energy threshold to the imaging scan and subsequent data processing.

[0063] In another aspect, the additional energy threshold can also be optimized by estimating the patient's shape during a scout scan prior to the imaging scan and adjusting the additional energy threshold based on the simulated spectrum. This spectrum is simulated by calculating the shortest path length / attenuation at each view angle and fan angle together with the post-dose-compensation spectrum calculated as a function of the fan angle.

[0064] Depending on the embodiment, various groups of pixels or each pixel can have a different number of energy bins. For example, the central pixel group can have 5 energy bins, the peripheral pixel group can have 3 energy bins, and the intermediate pixel group can have 4 energy bins. Here, for example, the central pixel group is a group of pixels where the central X-ray 115 is incident, the peripheral pixel group is a group of pixels where the edge X-ray 105 is incident, and the intermediate pixel group is a group of pixels where the intermediate X-ray 110 is incident. That is, the number of energy bins set for each group of pixels is variable. When the first energy threshold of the pixel is set relatively high, if the size of the energy bins is made uniform and the energy bins are simply shifted in the higher energy direction, higher energy bins will be compressed along the energy axis.

[0065] For example, the first energy threshold of the edge spectrum 105a can be set to 40 keV, but this will cause the fifth energy bin 120e to shift and measure from 120 keV to 140 keV. Note that here the spectra do not contain the same level of information. Therefore, the fifth energy bin 120e is completely removed, leaving four energy bins 120. In another aspect, when the scan is set to scan a specific object of a patient, such as in the case of a heart scan, and the edge part of the patient contains low-value information, the energy bin 120 of the edge detector (the pixels in the edge part of the detector) is removed. In this way, one to two energy bins of the edge detector may be removed. If a certain pixel group has fewer energy bins than other pixel groups, the pixel group with fewer energy bins often has a higher value as the first energy threshold. Similarly, for example, the number of energy bins of a certain pixel (the first detection element) may be made less than the number of energy bins of another pixel (the second detection element), and the number of energy bins of the first detection element may be two or more. In this case, the first energy threshold of the first detection element may be made larger than the first energy threshold of the second detection element.

[0066] In summary, depending on the complexity of the scan settings, the operator can select the first energy threshold and additional energy thresholds based only on the predetermined filter used if they desire a faster scan. Similarly, the operator can have the CT system perform a series of scout scans to determine the first energy threshold for all groups of pixels regardless of the orientation of the radiation source if they desire a slower and more accurate scan. The operator may desire to perform an accurate scout scan for the slowest and most accurate scan. In this case, the first energy threshold can be adjusted for each view (view angle) while rotating the radiation source in all directions around the patient.

[0067] The above-described features regarding the reduction in the number of energy bins provide a particular advantage regarding the reduction in the amount of data transmitted via the slip ring of the CT system. Since the bandwidth of the slip ring is generally limited, such a reduction in the amount of data generated and transmitted is of great significance.

[0068] Also, the above-described features benefit a CT system in which the first energy threshold and the additional energy threshold are fixed. This is because the optimal first energy threshold and additional energy threshold for a group of pixels are determined based on beam filtering alone, or based on beam filtering and scout scan information. Thus, each of the energy bins then collects information from the imaging scan in the most effective and optimal manner.

[0069] FIG. 4 is a diagram showing an example of the configuration of a CT scanner (computed tomography scanner, X-ray CT apparatus) according to the present embodiment. The CT scanner is an example of an imaging apparatus. As shown in FIG. 4, the CT scanner includes an X-ray imaging gantry 500, an X-ray tube 501, an annular frame 502, and a multi-row (multi-column) or two-dimensional array type X-ray detector 503. The X-ray tube 501 and the X-ray detector 503 are mounted diametrically on the annular frame 502 with an object OBJ sandwiched therebetween. The annular frame 502 is supported so as to be rotatable around a rotation axis RA (i.e., the axis of rotation). The rotating device 507 rotates the annular frame 502 at a high speed such as 0.4 seconds / rotation, while the object OBJ moves along the axis RA to the back side or the front side in FIG. 4.

[0070] There are various types of X-ray CT devices. For example, there are rotate / rotate type devices in which both the X-ray tube and the X-ray detector rotate around the object to be examined, and stationary / rotate type devices in which many detection elements are arranged in a ring or planar shape and only the X-ray tube rotates around the object to be examined. The present disclosure can be applied to any type of device. Here, for the sake of simplicity, the case of using a rotate / rotate type device as an example will be described.

[0071] The multi-slice X-ray CT device further includes a high-voltage generator 509 that generates a tube voltage applied to the X-ray tube 501 via a slip ring 508 in order for the X-ray tube 501 to generate X-rays. X-rays are radiated onto an object OBJ having a cross-sectional area indicated by a circle. For example, the X-ray tube 501 has an average X-ray energy during a first scan that is smaller than the average X-ray energy during a second scan. Therefore, two or more scans can be obtained corresponding to different X-ray energies. The X-ray detector 503 is located on the opposite side of the X-ray tube 501 with the object OBJ therebetween in order to detect the X-rays that have passed through the object OBJ after being radiated. The X-ray detector 503 further includes individual detection elements or devices.

[0072] The CT apparatus further includes other apparatuses (such as a processing circuit) that process the signals detected by the X-ray detector 503. The data acquisition circuit or data acquisition system (Data Acquisition System: DAS) 504 converts the signal output of each channel from the X-ray detector 503 into a voltage signal, amplifies the signal, and further converts the signal into a digital signal. The X-ray detector 503 and the DAS 504 are configured to process a predetermined total number of projections per rotation (Total number of Projections Per Rotation: TPPR). The DAS 504 may include a comparator having a configurable energy threshold for discriminating the detected counts by energy bin. These energy levels can be set by various steps of the methods 200 and 300 of the present application.

[0073] The comparator compares the signal generated at the time of X-ray detection with a plurality of signal thresholds corresponding to each of the plurality of energy thresholds including the first energy threshold. The comparator determines that the X-ray has energy within the first energy bin range when the signal generated at the time of X-ray detection is greater than the first signal threshold and less than the second signal threshold.

[0074] The above-described data is sent to a preprocessing device (preprocessing circuit) 506 housed in a console outside the X-ray imaging gantry 500 via a non-contact data transmitter 505. The preprocessing device 506 performs various corrections such as sensitivity correction on the raw data. The storage device 512 stores the resultant data, also called projection data, at a stage immediately before the reconstruction process. The storage device 512 is connected to the system controller 510 via a data / control bus 511 together with a reconstruction device (reconstruction circuit) 514, an input device 515, and a display device 516. The system controller 510 controls a current regulator 513 that limits the current to a level sufficient to drive the CT system.

[0075] The detector is rotated and / or fixed in relation to the patient in various generations of CT scanner systems. In one embodiment, the CT system described above can be, for example, a system that combines a third-generation arrangement and a fourth-generation arrangement. In the third-generation system, the X-ray tube 501 and the X-ray detector 503 are mounted diametrically on the annular frame 502 and rotate around the object OBJ in response to the rotation of the annular frame 502 around the rotation axis RA. In the fourth-generation system, the detector is fixedly installed around the patient and the X-ray tube rotates around the patient. In other embodiments, the X-ray imaging gantry 500 includes a number of detectors arranged on an annular frame 502 supported by a C-arm and a stand.

[0076] The storage device 512 is realized by various memories. The storage device 512 can store measurement values representing the irradiation of X-rays on the X-ray detector 503. Further, the storage device 512 can store, for example, dedicated programs for executing various steps of methods 200 and 300 for reducing imaging artifacts.

[0077] The reconstruction device 514 can execute various steps of methods 200 and 300. Further, the reconstruction device 514 can execute pre-reconstruction processing including image processing such as necessary volume rendering processing and image difference processing.

[0078] The pre-reconstruction processing can include various steps of methods 200 and 300. Also, the pre-reconstruction processing of the projection data performed by the preprocessing device 506 can include, for example, detector calibration, detector non-linearity correction (non-linearity correction), and correction of the polarity effect.

[0079] The post-reconstruction processing performed by the reconstruction device 514 can include, if necessary, filtering and smoothing of the image, volume rendering processing, and image difference processing. The process of image reconstruction can execute various CT image reconstruction methods. The reconstruction device 514 can use, for example, a storage device (e.g., storage device 512) that stores projection data, reconstructed images, calibration data, parameters, and computer programs.

[0080] The preprocessing device 506, the system controller 510, and the reconstruction device 514 can include a CPU (processing circuit) that can be realized as individual logic gates such as an ASIC, a Field Programmable Gate Array (FPGA), or other Complex Programmable Logic Device (CPLD). The realization of the FPGA or CPLD can be coded in VHDL, Verilog, or other hardware description languages, and the code can be stored in the direct internal electronic memory of the FPGA or CPLD, or as separate electronic memory. Furthermore, the storage device 512 can be a non-volatile memory such as a ROM, EPROM, EEPROM, or FLASH memory. The storage device 512 can also be a volatile memory such as static or dynamic RAM, and a processing device such as a microcontroller or microprocessor can be provided to manage the interaction between the electronic storage device and the FPGA or CPLD and the storage device.

[0081] As an alternative, there is a method of executing a computer program including a series of computer-readable instructions for performing the functions described in the present application by the CPUs in the preprocessing device 506, the system controller 510, and the reconfiguration device 514. This program is stored in the above-described persistent electronic storage device and / or hard disk drive, CD, DVD, FLASH drive, or other known storage media. Further, the computer-readable instructions can be provided as a utility application, a background daemon, or a component of the operating system, or a combination thereof, that works with a processing device such as an Intel Xenon processor in the United States or an AMD Opteron processor in the United States, and an operating system such as Microsoft VISTA, UNIX (registered trademark), Solaris, LINUX (registered trademark), Apple, MAC-OS, and other operating systems well-known to those skilled in the art. Further, the CPU can be realized as a number of processing devices that cooperate and operate in parallel to execute the instructions.

[0082] In one embodiment, the reconstructed image can be displayed on the display device 516. The display device 516 can be an LCD display, a CRT display, a plasma display, an OLED, an LED, or other displays known in the art.

[0083] The storage device 512 can be a hard disk drive, a CD-ROM drive, a DVD drive, a FLASH drive, a RAM, a ROM, or other electronic storage devices known in the art.

[0084] PCD can use a direct X-ray irradiation detector that applies semiconductors such as cadmium telluride (CdTe), cadmium zinc telluride (CZT), silicon (Si), mercuric iodide (HgI2), and gallium arsenide (GaAs). A semiconductor-applied direct X-ray detector generally has a much faster time response than indirect detectors such as scintillator detectors. The rapid time response of the direct detector enables the resolution of individual X-ray detection events by this device. However, in a typical high X-ray beam in clinical X-ray applications, some retention of the detection phenomenon occurs. Since the energy of the detected X-ray is proportional to the signal generated by the direct detector, the detection events can be grouped by energy bin and the X-ray data of spectral CT can be spectrally decomposed.

[0085] Certain embodiments have been described, but these embodiments are presented for illustrative purposes only and are not intended to limit the teachings of this disclosure. In fact, the new methods, devices, and systems described in this application can be embodied in various other forms. Furthermore, various omissions, substitutions, and changes can be made to the forms of the methods, devices, and systems described herein without departing from the gist of this disclosure.

[0086] According to at least one of the embodiments described above, optimization of the energy bin can be achieved.

[0087] Some embodiments of the present invention have been described, but 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 forms, and various omissions, replacements, and changes can be made without departing from the gist of the invention. These embodiments and their modifications are included in the scope and gist of the invention, as well as in the invention described in the claims and its equivalent scope.

Description of Reference Numerals

[0088] 506 Pretreatment Device 514 Reconfiguration Device

Claims

Claim 1 For two or more detection elements of a photon counting detector, estimating the respective energy spectra of the X-ray beams of an X-ray source incident on the corresponding detection elements by modeling X-ray attenuation as a function of X-ray energy, An imaging apparatus comprising a processing circuit that, for each of the detection elements, sets, as the first energy threshold, the minimum energy of the X-rays incident on the detection element indicated by the energy spectrum estimated for the X-ray beam incident on the detection element, within the range of energies of the X-rays that are not counted from 0 keV to the first energy threshold. Claim 2 The processing circuit discriminates that the X-rays detected by one of the two or more detection elements are included within the first energy bin when the measured energy of the X-rays detected by the one detection element is less than a second energy threshold that is the maximum value of the energy range of the first energy bin of the one detection element and greater than the first energy threshold that is the minimum value of the energy range of the first energy bin of the one detection element. The imaging apparatus according to claim 1. Claim 3 The processing circuit reduces the number of energy bins of a first detection element among the two or more detection elements to be less than the number of energy bins of a second detection element among the two or more detection elements, and sets the number of the energy bins of the first detection element to be 2 or more. The imaging apparatus according to claim 1 or 2. Claim 4 The processing circuit makes the first energy threshold of the first detection element greater than the first energy threshold of the second detection element. The imaging apparatus according to claim 3. Claim 5 The processing circuit estimates the respective energy spectra by modeling X-ray attenuation when the X-ray beam passes through a filter. The imaging apparatus according to any one of claims 1 to 4. Claim 6 The processing circuit estimates the respective energy spectra by modeling X-ray attenuation when the X-ray beam passes through a filter and an object. The imaging apparatus according to any one of claims 1 to 4. Claim 7 The imaging device according to claim 6, wherein the processing circuit estimates each of the energy spectra by modeling the energy spectrum based on a simulation of the X-ray beam passing through the filter and the object.

8. The imaging device according to claim 6, wherein the processing circuit estimates each of the energy spectra by calibrating the X-ray beam passing through the filter and modeling the energy spectrum based on a simulation of the X-ray beam passing through the object.

9. The processing circuit estimates each of the energy spectra of the X-ray beam by determining a shortest path length, which is a length of a shortest path from a source of the X-ray beam passing through the object to the object at one view angle, and performs a simulation of the X-ray beam passing through the object at the one view angle using the shortest path length. The imaging device according to claim 6.

10. The imaging device according to claim 9, wherein the shortest path length at the one view angle is a shortest path length at one of the view angles among the shortest path lengths at a plurality of view angles.

11. The imaging device according to claim 2, wherein the processing circuit sets, for each of the two or more detection elements, additional energy thresholds that define boundaries of respective energy bins, the additional energy thresholds including a second energy threshold between the first energy bin and the second energy bin.

12. The processing circuit groups the detection elements into a plurality of groups, each of the plurality of groups having a respective energy spectrum of the X-ray beam, and sets the first energy threshold for each of the groups based on the energy spectrum corresponding to each of the groups. The imaging device according to claim 1 or 2.

13. The imaging device according to claim 11, wherein a spectral width of the energy bin is not uniform.

14. The processing circuit estimates each of the energy spectra of the X-ray beam by determining a shortest path length, which is a length of a shortest path from a source of the X-ray beam passing through the object to the object at each view angle, The imaging apparatus according to claim 6, wherein simulation of the X-ray beam that has passed through the object is performed at the corresponding view angle using the shortest path length at each of the view angles.

15. The imaging apparatus further includes a comparator that compares a signal generated at the time of X-ray detection with a plurality of signal thresholds corresponding to the respective energy thresholds including the first energy threshold, wherein the comparator determines that the X-ray has energy within the range of the first energy bin when the signal generated at the time of X-ray detection is greater than the first signal threshold and less than the second signal threshold. The imaging apparatus according to claim 2.

16. Estimating, for two or more detection elements of a photon counting detection element, the respective energy spectra of the X-ray beams of an X-ray source incident on the corresponding detection elements by modeling X-ray attenuation as a function of X-ray energy; For each of the detection elements, setting, as the first energy threshold, the minimum energy of the X-rays incident on the detection element indicated by the energy spectrum estimated for the X-ray beam incident on the detection element within the range of the energy of the X-rays not counted from 0 keV to the first energy threshold; A method comprising.

17. The step of estimating the respective energy spectra includes, by using a scout scan to obtain scout projection data using a dose lower than the dose used in a subsequent imaging scan, modeling the X-ray attenuation after the X-rays have passed through an object, The method according to claim 16, comprising a process of estimating the respective energy spectra.

18. For two or more detection elements of a photon counting detection element, a process of estimating the respective energy spectra of the X-ray beams of an X-ray source incident on the corresponding detection elements by modeling X-ray attenuation as a function of X-ray energy; For each of the detection elements, a process of setting, as the first energy threshold, the minimum energy of the X-rays incident on the detection element indicated by the energy spectrum estimated for the X-ray beam incident on the detection element within the range of the energy of the X-rays not counted from 0 keV to the first energy threshold; A program for causing a computer to execute.

Images