Measurement and correction of sparse background to improve image quality.

Abstract

Disclosed herein is an imaging system including a first X-ray source configured to generate first X-ray photons in a first energy range suitable for imaging and project the first X-ray photons onto a region designated for imaging, a rotatable gantry configured to rotate the first X-ray source so that the first X-ray source traverses an angular path, and a data processor having an analysis portion configured to collect first data regarding the transmission of the first X-ray photons through the region designated for imaging at a set of image collection angles along the angular path and collect background data at a set of background collection angles along the angular path, where the system obtains two or more images of the region designated for imaging between the background angles. The analysis portion is also configured to remove errors in the first data using the background data and generate a corrected image based on the removal of errors in the first data.

Description

[Technical Field]

[0001] Cross-reference of related applications This application relates to U.S. Patent Application No. 16 / 694,148, “APPARATUS AND METHODS FOR SCALABLE FIELD OF VIEW IMAGING USING A MULTI-SOURCE SYSTEM,” filed November 25, 2019, by D. Gagnon et al. This application also includes applications No. 62 / 773,712 (agent number 38935 / 04001) filed on November 30, 2018, No. 62 / 773,700 (agent number 38935 / 04002) filed on November 30, 2018, No. 62 / 796,831 (agent number 38935 / 04004) filed on January 25, 2019, No. 62 / 800,287 (agent number 38935 / 04003) filed on February 1, 2019, No. 62 / 801,260 (agent number 38935 / 04006) filed on February 5, 2019, and No. 62 / 813,335 (agent number 38935 / 04006) filed on March 4, 2019. This relates to 11 U.S. provisional patent applications, including No. 62 / 821,116 (Agent No. 38935 / 04009) filed on March 20, 2019, No. 62 / 836,357 (Agent No. 38935 / 04016) filed on April 19, 2019, No. 62 / 836,352 (Agent No. 38935 / 04017) filed on April 19, 2019, No. 62 / 843,796 (Agent No. 38935 / 04005) filed on May 6, 2019, and No. 62 / 878,364 (Agent No. 38935 / 04008) filed on July 25, 2019. This application also includes, "MULTIMODAL RADIATION APPARATUS AND METHODS" (Filing No. 16 / 694,145, filed November 25, 2019, Agent Reference Number 38935 / 04019), and "INTEGRATED HELICAL FAN-BEAM COMPUTED TOMOGRAPHY IN IMAGE-GUIDED RADIATION TREATMENT DEVICE" (Filing No. 16 / 694, filed November 25, 2019).Application No. 161 (Agent Reference Number 38935 / 04011), Application No. 16 / 694,166 (Agent Reference Number 38935 / 04010) filed on November 25, 2019, entitled "COMPUTED TOMOGRAPHY SYSTEM AND METHOD FOR IMAGE IMPROVEMENT USING PRIOR IMAGE", Application No. 16 / 694,177 (Agent Reference Number 38935 / 04013) filed on November 25, 2019, entitled "OPTIMIZED SCANNING METHODS AND TOMOGRAPHY SYSTEM USING REGION OF INTEREST DATA", "HELICAL CONE-BEAM COMPUTED TOMOGRAPHY IMAGING WITH AN OFF-CENTERED Application No. 16 / 694,190, filed November 25, 2019, titled "DETECTOR" (Agent reference number 38935 / 04015), Application No. 16 / 694,192, filed November 25, 2019, titled "MULTI-PASS COMPUTED TOMOGRAPHY SCANS FOR IMPROVED WORKFLOW AND PERFORMANCE" (Agent reference number 38935 / 04021), Application No. 16 / 694,202, filed November 25, 2019, titled "METHOD AND APPARATUS FOR SCATTER ESTIMATION IN CONE-BEAM COMPUTED TOMOGRAPHY" (Agent reference number 38935 / 04012), "ASYMMETRIC SCATTER FITTING FOR OPTIMAL PANEL READOUT IN CONE-BEAM COMPUTED Application No. 16 / 694,210 dated November 25, 2019, titled "TOMOGRAPHY" (Agent reference number 38935 / 04014), Application No. 16 / 694 dated November 25, 2019, titled "METHOD AND APPARATUS FOR IMPROVING SCATTER ESTIMATION AND CORRECTION IN IMAGING"Relates to ten non-provisional U.S. patent applications filed on the same day, including U.S. Patent Application No. 218 (Agent Docket No. 38935 / 04018) and U.S. Patent Application No. 16 / 694,230 (Agent Docket No. 38935 / 04022) filed on November 25, 2019, entitled "METHOD AND APPARATUS FOR IMAGE RECONSTRUCTION AND CORRECTION USING INTER-FRACTIONAL INFORMATION". All of the content of the patent applications and patents (if any) identified above are hereby incorporated by reference in their entirety into this specification.,

[0002] Aspects of the disclosed technology relate to patient imaging using X-ray radiation during, before, and after radiation therapy (RT). The disclosed technology relates in particular to removing residual effects caused by charging a particular X-ray detector. Further, it relates to measuring background radiation and using the background radiation to correct the detected image., BACKGROUND OF THE INVENTION

[0003] External beam radiation therapy is a non-invasive treatment alternative to high-risk and invasive surgery. It can use X-rays generated by a treatment radiation source such as a linear accelerator (LINAC) to treat pathological anatomical structures (e.g., tumors, lesions, angiogenesis abnormalities, neuropathies, etc.). Typically, the radiation source directs X-ray beams at the tumor site from multiple angles. By carefully controlling the orientation of the radiation source, it can be ensured that each X-ray beam passes through the same tumor site but through different regions of adjacent healthy tissue. This keeps the cumulative radiation dose high in the tumor while keeping the dose in healthy tissue relatively low.,

[0004] "Radiosurgery" refers to irradiating a target area with a dose sufficient to cause necrosis of the lesion more rapidly than radiotherapy. Radiosurgery involves higher single doses (e.g., 500–2000 centigrays) and hypofractionated irradiation (e.g., 1–5 sessions or treatment days). In contrast, radiotherapy may use 100–200 centigrays and multifractionated irradiation (e.g., 30–45 sessions). X-ray sources for radiotherapy and radiosurgery tend to be in the MeV range, which is a higher energy than imaging sources, which tend to be in the keV range. The terms "radiotherapy" and "radiotherapy" are used interchangeably herein to mean radiosurgery and / or radiotherapy for convenience in contrasting these two MeV techniques with imaging techniques using keV X-rays, unless otherwise specified. In this specification, "MV" and "MeV" X-rays and sources are referred to interchangeably, as is customary, with "kV" and "keV" X-rays and sources. When specifically referring to the energy of X-rays and sources, "MeV" and "keV" are used.

[0005] Image-guided radiotherapy (IGRT) systems combine keV and MeV sources for imaging and treatment. IGRT systems are typically classified by how the therapeutic X-ray source is mounted and moved. In gantry IGRT, the gantry rotates the therapeutic radiation source around an axis passing through the "isocenter" or intersection of the X-ray beam. The result is an X-ray beam intersection in a 3D volume of spherical or ellipsoidal shape. Different types of gantry mount and move the X-ray source differently. C-arm gantry mounts the therapeutic radiation source within a cantilever and rotates it around an axis passing through the isocenter. Ring gantry mounts the therapeutic source in an annular or ring-shaped element. The patient's body extends through a hole in the annular or ring. The annular or ring rotates around an axis passing through the isocenter. In robotic arm-based systems, the therapeutic radiation source is mounted on a robotic arm, increasing its degrees of freedom of movement. The robotic arm extends above and around the patient. This can provide at least five degrees of freedom for delivering therapeutic radiation from multiple out-of-plane directions. In contrast, a ring or C-arm system delivers therapeutic radiation with a set angle defined by the rotational trajectory of the radiation source.

[0006] X-ray imaging systems can be integrated into radiotherapy systems to guide radiation delivery. They can also track the movement of the target during treatment. MeV imaging systems may have detectors positioned opposite the treatment source to image the patient for setup and treatment imaging. Other approaches use separate, independent imaging radiation sources and / or detectors for patient setup and treatment imaging. Tracking the target during treatment becomes possible by comparing treatment images with previous or pre-treatment imaging information. Pre-treatment imaging information may include, for example, CT data, cone-beam CT (CBCT) data, magnetic resonance imaging (MRI) data, positron emission tomography (PET) data, or 3D rotational angiography (3DRA) data, and any information obtained from these imaging modalities (e.g., digitally reconstructed radiographs (DRRs), but not limited to these).

[0007] As explained above, keVX sources are typically used for imaging. KeV sources tend to provide good contrast for most types of soft tissue. However, keV systems do not perform well when the imaging zone contains dense tissue (e.g., thick bone, calcified arteries, etc.). If the tissue of interest is located within the same irradiation area as these dense materials, the keV image may be deficient due to the interaction between the keVX-rays and the dense material. Streak artifacts can obscure or darken the area of ​​interest. Metals along the X-ray path (e.g., dental fillings, implants, or stents) can cause photon starvation, obscuring the area of ​​interest. Scattering errors and other issues can lead to additional problems. Another important issue is "beam hardening," where certain parts of the body "see" different X-ray photon spectra due to selective absorption in other parts of the body.

[0008] MeVX sources, primarily used in treatment, can also generate images. However, the contrast-to-noise ratio (CNR) in MeV images can be low. Some tissues have good intrinsic contrast in the keV range, while others have good contrast in MeV. However, at a fitted dose, which is a crucial factor in patient imaging, there are fewer MeV X-rays. Furthermore, MeV X-rays are more difficult to detect, and these measurements are noisier. This generally worsens the MeVCNR for all tissues.

[0009] Crude MeV X-ray images are often used to determine the position of the MeV therapeutic beam relative to the patient. However, there is untapped potential for more ambitious MeV imaging applications. In particular, combining keV and MeV X-ray imaging in a single therapeutic device can utilize the complementary advantages of both energy ranges. For example, MeV X-ray contrast can fill gaps in keV images where denser tissue degrades the keV image contrast. Therefore, improved and enhanced techniques for combining information from both types of images are disclosed below.

[0010] Each of the systems described above can present problems in image acquisition and analysis. For example, X-rays used in keV imaging can transiently affect the detector, and this can persist for two or more image scans. This can result in abnormal features appearing in images obtained from subsequent scans, which can mask image features and complicate analysis. Such effects can be observed with high-energy MeVX-rays used in radiotherapy, due to the use of certain detectors that are prone to charging and other residual effects. [Overview of the project]

[0011] According to aspects of the present disclosure, the imaging system includes a first X-ray source configured to generate first X-ray photons in a first energy range suitable for imaging and to project the first X-ray photons onto a region designated for imaging; a rotatable gantry configured to rotate the first X-ray source so that the first X-ray source traverses an angular path; and a data processor having an analysis unit. The analysis unit is configured to collect first data relating to the transmission of first X-ray photons through the region designated for imaging in a set of image acquisition angles along the angular path, and to collect background data in a set of background acquisition angles along the angular path, such that the number of image acquisition angles is maximized among the background acquisition angles. The analysis unit is also configured to use the background data to remove errors from the first data and to generate a corrected image based on the removal of errors from the first data.

[0012] The analysis part can determine, at least partially, a set of background acquisition angles by determining the minimum number of background measurements. The analysis part can determine, at least partially, a set of background acquisition angles by maximizing the set of image acquisition angles between background acquisition angles. Removing errors from the first data using background data may include weighting the image using the measured time decay response of the X-ray detector and subtracting the weighted image from the first data. The analysis part may recalibrate the time decay response based on background data. Recalibrating the time decay response may be based on two or more background images and include two or more decay terms. Removing errors from the first data using background data may include interpolating the background images.

[0013] The modified form includes an X-ray source configured to generate X-ray photons in an energy range suitable for imaging and to project the X-ray photons onto a region designated for imaging, and a rotatable gantry configured to rotate the X-ray source so that the first X-ray source traverses an angular path. The system includes a data processor having an analysis unit configured to collect background data at a first angle along an angular path, collect image data relating to the transmission of the X-ray photons through the region designated for imaging over an angular range along the angular path, and collect background data at a second angle along the angular path, wherein the second angle can be positioned such that the angular range is between the first and second angles, and generate a corrected image by removing errors in the image data using interpolation of the background data collected at the first and second angles.

[0014] The analysis unit may, at least partially, determine the first angle, second angle, and angular range by determining the minimum number of background measurements for accurate background estimation. The analysis unit may, at least partially, determine the first angle, second angle, and angular range by maximizing at least one of the total number of images of the region designated for imaging and the number of images of the region designated for imaging taken between each background acquisition angle. Removing errors from the image data may include weighting the images using the measured time decay response of the X-ray detector and subtracting the weighted images from the first data. The analysis unit may recalibrate the time decay response based on the background data. This system includes a cone-beam computed tomography (CBCT) system. The system may include a second X-ray source configured to generate second X-ray photons in a second energy range with different energies from the first energy range, and to project the second X-ray photons onto a region designated for imaging, wherein the analysis unit may be configured to combine data obtained from the first and second X-ray photons, interleave the irradiation of the first and second X-ray sources, and operate the first and second X-ray sources simultaneously. The set of background acquisition angles may be evenly distributed along the angular path. The set of background acquisition angles may be distributed at predetermined angles along the angular path. The set of background acquisition angles may be determined through at least one of the following: analysis of planned computed tomography (CT) images, analysis of previously available CT images, analysis of CT2D survey images, analysis of orthogonal angle images, analysis of patient atlases, and analysis of data being acquired.

[0015] The set of background acquisition angles may be determined using at least one of the patient's geometric shape, acquisition protocol, and projection angle relative to the patient's orientation. The analysis portion may determine the set of background acquisition angles via an algorithm. The algorithm may determine the set of background acquisition angles based on improving the accuracy of CBCT reconstruction. The algorithm may exclude conjugate angles from the set of background acquisition angles. The algorithm may select background acquisition angles based at least in part on the use of MeV radiation. The algorithm may select background acquisition angles to improve at least one of lag correction and data-driven scattering correction. Removing errors in the first data using background data may include: acquiring background data at a first angle along an angular path; acquiring first data within an angular range along an angular path; acquiring background data at a second angle along an angular path, wherein the second angle is positioned such that the angular range can be between the first and second angles; generating an error image by interpolating the background data acquired at the first and third angles; and removing errors in the first data using the error image.

[0016] The algorithm can generate a first set of background acquisition angles for a first scan and a second set of background acquisition angles for a second scan, and the first and second sets may be different. The first and second sets may not overlap. The angular path may be helical. The algorithm may determine the background acquisition angles at least partially based on the pitch size of the helix. The algorithm may determine the background acquisition angles at least partially based on improving the accuracy of 3D image reconstruction using image data. Image reconstruction may select different sets of background acquisition angles for different rotations of the first source. The analysis unit can synchronize X-ray control, flat panel detector readout, and CT scan control so that when the power to the first X-ray source is turned off, the flat panel detector readout provides background data. The movement of the first X-ray source along the angular path may not be interrupted while the X-ray source is turned off. The analysis unit may be further configured to generate 3D reconstructions by modifying the reconstruction algorithm based on the background acquisition angles.

[0017] The correction takes into account the non-uniform intervals between background acquisition angles and the non-uniform intervals between image acquisition angles. The correction takes into account the difference between the azimuth position of the background image in the first rotation of the first X-ray source and the azimuth position of the background image in the second rotation of the first X-ray source. The azimuth position of the first rotation can be interleaved with the azimuth position of the second rotation. A method for operating an imaging system, comprising: generating a first X-ray photon in a first energy range suitable for imaging; projecting the first X-ray photon onto a region designated for imaging; rotating the first X-ray source so that the first X-ray source traverses an angular path; collecting first data relating to the transmission of the first X-ray photon through the region designated for imaging at a set of image acquisition angles along the angular path; collecting background data at a set of background acquisition angles along the angular path, wherein the system acquires two or more images of the region designated for imaging between background angles; removing errors from the first data using the background data; and generating a corrected image based on the removal of errors from the first data.

[0018] The modified form includes an imaging system comprising: a first X-ray source configured to generate first X-ray photons in a first energy range suitable for imaging and to project the first X-ray photons onto a region designated for imaging; a rotatable gantry configured to rotate the first X-ray source so that the first X-ray source traverses an angular path; and a data processor having an analysis section configured to collect first data relating to the transmission of first X-ray photons through a region designated for imaging at a set of image acquisition angles along the angular path, collect background data at a background acquisition angle along the angular path, remove errors from the first data using the background data, and generate a corrected image based on the removal of errors from the first data.

[0019] Features described and / or illustrated in relation to one embodiment may be used in the same way, similarly, and / or in combination with or instead of features of other embodiments in one or more other embodiments.

[0020] This description of the present invention does not limit in any way the terms used in the claims, nor does it limit the scope of the claims or the present invention. The terms used in the claims have all their ordinary meanings.

[0021] The accompanying drawings incorporated herein and constituting part thereof illustrate embodiments of the present invention, which, together with the general description of the invention given above and the detailed description given below, are useful in illustrating embodiments of the invention. It will be understood that the element boundaries illustrated in the figures (e.g., boxes, groups of boxes, or other shapes) represent one embodiment of the boundary. In some embodiments, one element may be designed as multiple elements, or multiple elements may be designed as one element. In some embodiments, an element shown as an internal component of another element may be implemented as an external component, and vice versa. Furthermore, elements may not be drawn to scale. [Brief explanation of the drawing]

[0022] [Figure 1] This is a perspective view of an exemplary multimodal radiotherapy device according to one aspect of the disclosed technology. [Figure 2A] This is a schematic diagram of an exemplary multimodal radiotherapy apparatus according to one aspect of the disclosed technology. [Figure 2B] The rotation path 202 for the radiation source 30 and the rotation path 204 for the detector 34 within the gantry 12 are shown. [Figure 2C] The image shown was taken in an energy-concentrated region 206 that causes a charging or ghosting effect on subsequent images. [Figure 2D]It is the intensity profile of the image of FIG. 2C showing the energy concentration region 206. [Figure 2E] It is an image taken following the image of FIG. 2C having a residual or ghost generation effect 208 caused by the high energy region 206 of FIG. 2C. [Figure 2F] It is the profile of the image of FIG. 2E showing the residual or ghost generation effect 208. [Figure 3A] It shows the individual angular portions of the rotation path 202 where the background B or the image data I can be measured. [Figure 3B] It shows the path 202 for two different consecutive rotations 350 and 360 having alternating background B measurements. [Figure 4A] It is a flowchart showing a method 400 that can be implemented in accordance with the present disclosure. [Figure 4B] It is the detail of step 440 of the method 400.

Mode for Carrying Out the Invention

[0023] The following includes definitions of exemplary terms that can be used throughout the present disclosure. Both the singular and plural forms of all terms are included in each meaning.

[0024] As used herein, a "component" can be defined as a part of hardware, a part of software, or a combination thereof. A part of hardware can include at least a processor and a part of memory, and the memory includes instructions to execute. A component can be associated with a device.

[0025] As used herein, “logic” is synonymous with “circuit” and includes, but is not limited to, hardware, firmware, software, and / or any combination of functions or actions that perform those functions. For example, depending on the desired application or needs, logic may include software-controlled microprocessors, discrete logic such as application-specific integrated circuits (ASICs), or other programmed logic devices and / or controllers. Logic may also be fully embodied as software.

[0026] As used herein, “processor” includes, but is not limited to, any combination of substantially any number of processor systems or standalone processors, such as microprocessors, microcontrollers, central processing units (CPUs), and digital signal processors (DSPs). A processor may be associated with various other circuits that support the operation of the processor, such as random access memory (RAM), read-only memory (ROM), programmable read-only memory (PROM), erasable programmable read-only memory (EPROM), clocks, decoders, memory controllers, or interrupt controllers. These support circuits may be located inside or outside the processor or its associated electronic package. The support circuits communicate operationally with the processor. Support circuits are not necessarily shown separately from the processor in block diagrams or other drawings.

[0027] As used herein, “signal” includes, but is not limited to, one or more electrical signals, including analog or digital signals, one or more computer instructions, bits or bitstreams, etc.

[0028] "Software," as used herein, includes, but is not limited to, one or more computer-readable and / or executable instructions that cause a computer, processor, logic, and / or other electronic device to perform functions, actions, and / or behave in a desired manner. Instructions can be embodied in various forms, such as routines, algorithms, modules, or programs, which may include separate applications or code from dynamically linked sources or libraries.

[0029] While the above exemplary definitions are provided, the applicant intends that the broadest and most reasonable interpretation consistent with this specification be used for these terms and other terms.

[0030] As will be described in more detail below, embodiments of the disclosed technology relate to multimodal imaging / radiotherapy devices and methods. In some embodiments, radiotherapy delivery devices and methods may utilize, in conjunction with or as part of IGRT, an integrated low-energy radiation source for imaging and a high-energy radiation source for treatment and / or imaging. In particular, for example, radiotherapy delivery devices and methods may combine a low-energy collimated radiation source for imaging within a gantry with a high-energy radiation source for imaging and / or therapeutic treatment using rotational imaging acquisition (e.g., helical or step-and-shoot rotation with or without the ability to rotate continuously using slip rings).

[0031] Complementary information and advantages can be fully derived from keV and MeV radiation sources. For example, the inherent contrast of soft tissue may be higher at lower energies, but at higher energies, primary photons will not be starved through broad or dense structures. KeV and MeV imaging data can be used and complemented by each other to produce higher quality images. High-quality volumetric imaging may be required for visualization of target and risk organs (OARS), adaptive therapy monitoring, and treatment planning / replanning. In some embodiments, multimodal systems can also be used for positioning, motion tracking, and / or characterization or correction functions.

[0032] This image acquisition methodology may include, or otherwise utilize, multiple rotational scans, such as continuous scans (e.g., longitudinal movement of the patient support through the gantry bore, in addition to a helical source trajectory around the central axis), discontinuous circular stop-and-reverse scans with longitudinal incremental movement of the patient support, or step-and-shoot circular scans.

[0033] According to various embodiments, a multimodal apparatus uses a beamformer (which may include a collimator) to collimate a radiation source into, for example, a cone beam or a fan beam, in order to limit the beam. In one embodiment, the collimated beam can be combined with a gantry that rotates continuously while the patient moves, resulting in helical imaging.

[0034] In some embodiments, the time associated with increasing scan rotation to complete high-quality volumetric images can be mitigated by high gantry rate / speed (e.g., using high-speed slip-ring rotations including up to 10, 20, 60, or more rpm in revolutions per minute (rpm)), high frame rate, and / or sparse data reconstruction techniques to provide CT-quality imaging on a radiotherapy delivery platform. Detectors (having various row / slice sizes, configurations, dynamic ranges, etc.), scan pitch, and / or dynamic collimation are additional features in various embodiments, including selectively exposing portions of the detector and selectively defining active readout regions.

[0035] Multimodal apparatuses and methods can provide selective and variable collimation of a radiation beam emitted by a radiation source, including adjusting the radiation beam shape to expose a smaller area than the entire active area of ​​the associated radiation detector (e.g., a radiation detector positioned to receive radiation from an X-ray radiation source). Alternatively, by directly exposing only the primary region of the detector to radiation, the shadowed region of the detector can receive only scattered radiation. In some embodiments, scattering measurements in the shadowed region of the detector (and in some embodiments, measurements in the penumbra region) can be used to estimate scattering in the primary region of the detector receiving projection data.

[0036] Multimodal apparatuses and methods can provide selective and variable detector readout regions and ranges, including adjusting the detector readout range to limit the detector's active region in order to improve readout speed. For example, data less than the available shadow region data may be read and used for scattering estimation. The combination of selective readout and beamforming enables various optimizations of scattering fitting techniques.

[0037] Several exemplary embodiments of apparatus capable of carrying out the disclosed embodiments are described below. It should be understood that the embodiments are not limited to the specific hardware and apparatus disclosed herein. For example, any of the methods and algorithms disclosed herein may be carried out by the apparatus disclosed in U.S. Patent Application No. 16 / 694,148, filed November 25, 2019, which is incorporated herein in its entirety by reference.

[0038] Referring to Figures 1 and 2A, a multimodal apparatus 10 is shown. It will be understood that the multimodal apparatus 10 may be associated with and / or integrated with a radiotherapy device (such as shown in Figure 2A) that can be used for a variety of applications, including but not limited to an IGRT, such as an IGRT delivery system (e.g., an IGRT delivery system 104 shown in Figure 3A and discussed in detail below). The multimodal apparatus 10 includes a rotatable gantry system called a gantry 12, which is supported by or otherwise housed in a support unit or housing 14. In this specification, a gantry refers to a gantry system comprising one or more gantry (e.g., a ring or C-arm) capable of supporting one or more radiation sources and / or associated detectors as they rotate around a target. For example, in one embodiment, a first radiation source and its associated detector may be mounted on a first gantry of the gantry system, and a second radiation source and its associated detector may be mounted on a second gantry of the gantry system. In another embodiment, two or more radiation sources and associated detectors may be mounted on the same gantry of the gantry system, for example, when the gantry system consists of only one gantry. Various combinations of gantry, radiation sources, and radiation detectors may be combined into various gantry system configurations to image and / or treat the same volume within the same apparatus. For example, keV and MeV radiation sources can be mounted on the same or different gantry of the gantry system and used selectively for imaging and / or treatment as part of an IGRT system. When mounted on different gantry, the radiation sources can rotate independently but can still image the same (or nearly the same) volume simultaneously. The rotatable ring gantry 12 may have a capability of 10 rpm or more, as described above. The rotatable gantry 12 can define a gantry bore 16 and move and position a patient through this bore for imaging and / or treatment.According to one embodiment, the rotatable gantry 12 is configured as a slip-ring gantry that provides continuous rotation of the radiation source and associated radiation detector(s) while providing sufficient bandwidth for high-quality imaging data received by the detector(s). The slip-ring gantry eliminates alternating gantry rotations to wind up and unwind cables carrying power and signals associated with the device. Such a configuration enables continuous helical computed tomography, including CBCT, even when integrated into an IGRT system. As mentioned above, the main problem with single-rotation CBCT is insufficient sampling in all slices except the central slice (the slice including rotation). This can be overcome by helical trajectory cone-beam imaging.

[0039] The patient support unit 18 is positioned adjacent to the rotatable gantry 12 and is configured to support the patient in a typically horizontal position for longitudinal movement to and within the rotatable gantry 12. The patient support unit 18 can move the patient, for example, in a direction perpendicular to the plane of rotation of the gantry 12 (along or parallel to the axis of rotation of the gantry 12). The patient support unit 18 may be operably coupled to a patient support unit controller to control the movement of the patient and the patient support unit 18. The patient support unit controller can be synchronized with the rotatable gantry 12 and with a radiation source attached to the rotatable gantry that rotates around the patient's longitudinal axis according to a commanded imaging and / or treatment plan. Once the patient support unit is placed in the bore 16, it can also be moved up and down and left and right within a limited range to adjust the patient's position for optimal treatment. The axes x, y, and z (viewed from the front of gantry 12, the x-axis is horizontal and points to the right, the y-axis points to the gantry plane, and the z-axis is vertical and points upward) are shown. The x-axis, y-axis, and z-axis follow the right-hand rule.

[0040] As shown in Figure 2A, the multimodal apparatus 10 includes a low-energy radiation source (e.g., keV) 30 coupled to a rotating gantry 12 or otherwise supported by the rotatable gantry 12. In this embodiment, the low-energy radiation source 30 is an imaging radiation source that emits a radiation beam (generally indicated as 32) for producing high-quality images. In this embodiment, the imaging radiation source is an X-ray source 30 configured as a kiloelectron volt (keV) source (e.g., a clinical X-ray source having a voltage in the range of about 20 keV to about 150 keV). In one embodiment, the keV source of radiation includes kiloelectron volt peak photon energies (keV) up to 150 keV. The imaging radiation source can be any type of transmission source suitable for imaging. For example, the imaging radiation source may be, for example, an X-ray source (including for CT) or any other method of generating photons with sufficient energy and flux (e.g., a gamma-ray source (e.g., cobalt-57, with an energy peak at 122 keV), an X-ray fluorescence source (e.g., a Pb k-ray fluorescence source with two peaks at about 70 keV and about 82 keV), etc.). References to X-rays, X-ray imaging, X-ray imaging sources, etc., herein are illustrative for specific embodiments. Other imaging transmission sources may be used interchangeably in various other embodiments. The X-ray detector 34 (e.g., a two-dimensional planar detector or a curved detector) may be coupled to the rotatable gantry 12 or otherwise supported by the rotatable gantry 12. The X-ray detector 34 is positioned to receive radiation from the X-ray source 30 and can rotate with the X-ray source 30.

[0041] It will be understood that the X-ray detector 34 can take on many configurations without departing from the scope of the disclosed art. As shown in Figure 2A, the X-ray detector 34 can be configured as a flat panel detector (e.g., a multi-row flat panel detector). According to another exemplary embodiment, the X-ray detector 34 can be configured as a curved detector. The detector 34 can detect, or otherwise measure, the amount of unattenuated radiation and thus estimate (by comparison with the initially generated amount) the amount actually attenuated by the patient or associated patient ROI. The detector 34 can detect, or otherwise collect, attenuation data from different angles as the low-energy radiation source 30 rotates around the patient and emits radiation toward the patient.

[0042] Figures 1 and 2 show a multimodal apparatus 10 in which a radiation source 30 is mounted on a ring gantry 12, but other embodiments may include other types of rotatable imaging devices, such as C-arm gantry and robotic arm-based systems. In a gantry-based system, the gantry rotates the imaging radiation source 30 around an axis passing through an isocenter. The gantry-based system includes a C-arm gantry, in which the imaging radiation source 30 is mounted in a cantilever-like manner on an axis passing through an isocenter and rotates around that axis. The gantry-based system further includes a generally toroidal ring gantry, such as a rotatable gantry 12, through which the patient's body extends via a ring / toroidal bore, and the imaging radiation source 30 is mounted around the ring and rotates around an axis passing through an isocenter. In some embodiments, the gantry 12 rotates continuously. In other embodiments, the gantry 12 utilizes a cable-based system that repeatedly rotates and reverses.

[0043] A collimator or beamformer assembly (generally shown as 36) is positioned relative to the X-ray source 30 to selectively control and adjust the shape of the radiation beam 32 emitted by the X-ray source 30, thereby selectively exposing a portion or region of the active area of ​​the X-ray detector 34. The beamformer can also control how the radiation beam 32 is positioned on the X-ray detector 34. In one embodiment, the beamformer 36 may have one degree / dimensional movement (e.g., to create thinner or thicker slits). In another embodiment, the beamformer 36 may have two degrees / dimensional movement (e.g., to create rectangles of various sizes). In yet another embodiment, the beamformer 36 may be able to take on a variety of other dynamically controlled shapes, including, for example, parallelograms. All of these shapes can be dynamically adjusted during scanning. In some embodiments, the shielding portion of the beamformer can be rotated and / or translated.

[0044] The beam shaping unit 36 ​​can be controlled to dynamically adjust the shape of the radiation beam 32 emitted by the X-ray source 30 to several geometries, including, but not limited to, a fan beam or cone beam having a beam thickness (width) as small as the width of a single detector row, or multiple detector rows that can be only a portion of the detector's active area. In various embodiments, the beam thickness can expose a large detector active area of ​​several centimeters. For example, 3-4 centimeters (measured longitudinally in the detector plane) of a 5-6 centimeter detector can be selectively exposed to the imaging radiation 32. In this embodiment, 3-4 centimeters of projected image data can be acquired in each readout, leaving approximately 1-2 centimeters of unexposed detector area on one or both sides, which can be used to acquire scattering data, as described later.

[0045] In other embodiments, some portion of the active detector may be selectively exposed to imaging radiation. For example, in some embodiments, the beam thickness may be reduced to a range of similar sizes including detectors of about 2 cm, 1 cm, less than 1 cm, or smaller. In other embodiments, the beam thickness may be increased to a range of similar sizes including detectors of about 4 cm, 5 cm, greater than 5 cm, or larger. In various embodiments, the ratio of exposed detector area to active detector area may be 30-90% or 50-75%. In other embodiments, the ratio of exposed detector area to active detector area may be 60-70%. However, in other embodiments, various other sizes of exposed and active regions, or ratios of exposed detector area to active detector area, may be preferred. The beam and detector may be configured such that the shaded area of ​​the detector (active but not directly exposed to radiation) is sufficient to capture scattering data beyond the penumbra.

[0046] Various embodiments may include optimizing features that control the selective exposure of the detector (e.g., beam size, beam / aperture center, collimation, pitch, detector readout range, detector readout center, etc.) so that the measured data is sufficient for the primary (exposed) and shadow regions, but is also optimized for velocity and radiation dose control. The shape / position of the beam shaping unit 36 ​​and the readout range of the detector 34 can be controlled so that the radiation beam 32 from the X-ray source 30 covers a large or small portion of the X-ray detector 34 based on the specific imaging task and scattering estimation process being performed, including, for example, a combination of narrow and wide FOV scans. The apparatus 10 has the ability to acquire both single-rotation cone-beam images, as well as cone-beam images of helical or other wide and narrow beam angles.

[0047] The beam shaping unit 36 ​​can be configured in various ways that allow for adjustment of the shape of the radiation beam 32 emitted by the X-ray source 30. For example, the beam shaping unit 36 ​​may be configured to include a set of jaws or other suitable members that define and selectively adjust the size of the aperture through which the radiation beam from the X-ray source 30 can pass so as to be collimated. According to one exemplary configuration, the beam shaping unit 36 ​​may include an upper jaw and a lower jaw, which are movable in different directions (e.g., parallel directions) to adjust the size of the aperture through which the radiation beam from the X-ray source 30 passes, and to adjust the position of the beam 32 relative to the patient so as to optimize imaging by irradiating only the portion of the patient to be imaged and to minimize the patient dose.

[0048] According to one embodiment, the shape of the radiation beam 32 from the X-ray source 30 can be changed during image acquisition. In other words, according to one exemplary embodiment, the leaf position and / or aperture width of the beam shaping unit 36 ​​can be adjusted before or during the scan. For example, according to one embodiment, the beam shaping unit 36 ​​may be selectively controlled and dynamically adjusted so that, during rotation of the X-ray source 30, the radiation beam 32 has a shape that has a sufficient primary / shadow region and is adjusted to include only the target object (e.g., the prostate) during imaging. The shape of the radiation beam 32 emitted by the X-ray source 30 can be changed during or after the scan in accordance with the desired image acquisition, which can be based on imaging and / or therapeutic feedback, as will be discussed in more detail below.

[0049] As shown in Figure 2A, the multimodal apparatus 10 can be integrated with a radiotherapy device including a high-energy radiation source (e.g., MeV) 20 coupled to or otherwise supported by the rotatable gantry 12. According to one embodiment, the high-energy radiation source 20 is configured as a therapeutic radiation source, such as a high-energy radiation source used to treat a tumor in a patient in a region of interest. In other embodiments, the high-energy radiation source 20 is also configured as an imaging radiation source, or is used as at least such. It will be understood that the therapeutic radiation source may be a high-energy X-ray beam (e.g., a MeV X-ray beam) and / or a high-energy particle beam (e.g., an electron beam, a proton beam, or a beam of heavier ions such as carbon) or another preferred form of high-energy radiation. In one embodiment, the high-energy radiation source 20 includes a peak photon energy (MeV) of megaelectron volts of 1 MeV or more. In one embodiment, the high-energy X-ray beam has an average energy greater than 0.8 MeV. In another embodiment, the high-energy X-ray beam has an average energy greater than 0.2 MeV. In another embodiment, the high-energy X-ray beam has an average energy greater than 150 keV. Generally, the high-energy radiation source 20 has a higher energy level (peak and / or average, etc.) than the low-energy radiation source 30.

[0050] In one embodiment, the high-energy radiation source 20 is a LINAC that generates therapeutic radiation (e.g., MeV), and the imaging system includes a separate low-energy radiation source 30 that generates imaging radiation (e.g., keV) of relatively low intensity and lower energy. In other embodiments, the therapeutic radiation source 20 can be a radioisotope such as Co-60, which can generally have an energy greater than 1 MeV. The high-energy radiation source 20 can emit one or more radiation beams (generally shown as 22) toward a region of interest (ROI) within the patient supported on the patient support unit 18, according to the treatment plan.

[0051] In various embodiments, the high-energy radiation source 20 is used as both a therapeutic radiation source and an imaging radiation source. As will be described in detail below, the radiation sources 20 and 30 may be used in conjunction with each other to provide higher quality and better usable images. The reference to the therapeutic radiation source 20 is to distinguish the high-energy radiation source 20 from the low-energy radiation source 30, which may be used solely for imaging. However, the reference to the therapeutic radiation source 20 includes embodiments in which the therapeutic radiation source 20 (high-energy radiation source) may be used for therapy and / or imaging. In other embodiments, at least one additional radiation source may be coupled to the rotatable gantry 12 and operated to acquire projection data at a peak photon energy different from the peak photon energies of the radiation sources 20 and 30.

[0052] The detector 24 may be coupled to the rotatable gantry 12 or otherwise supported by the rotatable gantry 12 and positioned to receive radiation 22 from the therapeutic radiation source 20. The detector 24 may detect, or otherwise measure, the amount of unattenuated radiation and thus be able to estimate (by comparison with the initially generated amount) the amount actually attenuated by the patient or associated patient ROI. The detector 24 may detect, or otherwise collect, attenuation data from different angles as the therapeutic radiation source 20 rotates around the patient and emits radiation toward the patient.

[0053] It will be further understood that the therapeutic radiation source 20 may include a beam shaping device or collimator, or may be associated with it in other ways. A beam shaping device associated with the therapeutic radiation source 20 can be configured in several ways, similar to a beam shaping device 36 associated with the imaging source 30. For example, the beam shaping device may be configured as a multi-leaf collimator (MLC) that can include a plurality of combined leaves that are operable to move to one or more positions between a minimum open or closed position and a maximum open position. It will be understood that the leaves can be moved to desired positions to obtain a desired shape of the radiation beam emitted by the radiation source. In one embodiment, the MLC is capable of sub-millimeter targeting accuracy.

[0054] The therapeutic radiation source 20 can be mounted, configured, and / or moved in the same plane as the imaging source 30 or in a different plane (offset). In some embodiments, scattering caused by the simultaneous activation of radiation sources 20 and 30 can be incrementally reduced by offsetting the radiating planes. In other embodiments, scattering can be avoided by activating them alternately. For example, in simultaneous multimodal imaging, acquisitions can be performed in parallel rather than acquiring individual pulses simultaneously. In another embodiment, for example, to address the problem of MeV scattering on a keV detector, the use of shadow-based scattering correction can be used.

[0055] When integrated with a radiotherapy device, the multimodal device 10 can provide images used to set up (e.g., alignment and / or registration), plan, and / or guide a radiation delivery procedure (treatment). A typical setup is achieved by comparing the current (in-treatment) image with pre-treatment image information. Pre-treatment image information may include, for example, CT data, cone-beam CT data, MRI data, PET data, or 3D rotational angiography (3DRA) data, and / or any information obtained from these or other imaging modalities. In some embodiments, the multimodal device 10 can track the movement of the patient, target, or ROI during treatment.

[0056] The reconstruction processor 40 may be operably coupled to the detectors 24 and / or 34. In one embodiment, the reconstruction processor 40 is configured to generate a patient image based on radiation received by the detectors 24 and 34 from the radiation sources 20 and 30. It will be understood that the reconstruction processor 40 may be configured to be used to perform a method described more fully below. The apparatus 10 may also include a memory 44 suitable for storing information including, but not limited to, processing and reconstruction algorithms and software, imaging parameters, image data from previous or otherwise acquired images (e.g., planning images), treatment plans, etc.

[0057] The multimodal device 10 may include an operator / user interface 48, which allows the operator of the device 10 to interact with the device 10 or otherwise control the device 10 in order to provide input regarding scan or imaging parameters, etc. The operator interface 48 may include any suitable input device such as a keyboard, mouse, or voice-activated controller. The device 10 may also include a display 52 or other human-readable elements to provide output to the operator of the device 10. For example, the display 52 may allow the operator to view a reconstructed patient image and other information such as imaging or scan parameters related to the operation of the device 10.

[0058] As shown in Figure 2A, the multimodal apparatus 10 includes a controller (generally referred to as 60) operably coupled to one or more components of the apparatus 10. The controller 60 controls the overall function and operation of the apparatus 10, including supplying power and timing signals to the X-ray source 30 and / or therapeutic radiation source 20, and to the gantry motor controller which controls the rotational speed and position of the rotatable gantry 12. It will be understood that the controller 60 may include one or more of the following: a patient support controller, a gantry controller, a controller coupled to the therapeutic radiation source 20 and / or X-ray source 30, a beam shaping controller, a controller coupled to the detector 24 and / or X-ray detector 34, etc. In one embodiment, the controller 60 is a system controller that can control other components, devices, and / or controllers.

[0059] In various embodiments, the reconfiguration processor 40, operator interface 48, display 52, controller 60, and / or other components may be combined with one or more components or devices.

[0060] The device 10 may include various components, logic, and software. In one embodiment, the controller 60 includes a processor, memory, and software. For example, but not limited to, a multimodal device and / or radiotherapy system may include various other devices and components (e.g., gantry, radiation source, collimator, detector, controller, power supply, patient support, etc.) that can perform one or more routines or steps related to imaging and / or IGRT for a particular application. Routines may include imaging, image-based pre-delivery steps, and / or therapeutic delivery, including the respective device settings, configuration, and / or position (e.g., path / trajectory), which may be stored in memory. Furthermore, the controller(s) may directly or indirectly control one or more devices and / or components according to one or more routines or processes stored in memory. An example of direct control is setting various radiation source or collimator parameters (power, velocity, position, timing, modulation, etc.) related to imaging or treatment. An example of indirect control is communication of position, path, velocity, etc., to the patient support controller or other peripheral devices. The various controller hierarchies that may be associated with the device can be arranged in any suitable manner to communicate appropriate commands and / or information to the desired devices and components.

[0061] Furthermore, those skilled in the art will understand that the system and method can be implemented in other computer system configurations. Illustrated embodiments of the present invention can be implemented in a distributed computing environment in which specific tasks are performed by local or remote processing devices linked via a communication network. For example, in one embodiment, the reconfiguration processor 40 may be associated with a separate system. In a distributed computing environment, program modules may reside in both local and remote memory storage devices. For example, a remote database, a local database, a cloud computing platform, a cloud database, or a combination thereof can be used with the device 10.

[0062] The multimodal apparatus 10 can utilize exemplary environments for carrying out various aspects of the present invention, including a computer, which includes a controller 60 (for example, including a processor and memory 44) and a system bus. The system bus can connect system components, including but not limited to memory, to the processor and can communicate with other systems, controllers, components, devices, and the processor. The memory may include read-only memory (ROM), random access memory (RAM), hard drives, flash drives, and any other form of computer-readable media. The memory can store various software and data, including routines and parameters, which may include treatment plans.

[0063] The therapeutic radiation source 20 and / or X-ray source 30 can be operably coupled to a controller 60 configured to control the relative operation of the therapeutic radiation source 20 and the X-ray source 30. For example, the X-ray source 30 can be controlled and operated simultaneously with the therapeutic radiation source 20. In addition, or alternatively, the X-ray source 30 can be controlled and operated sequentially with the therapeutic radiation source 20, depending on the specific treatment and / or imaging plan being performed. For example, in various embodiments, the radiation sources 20 and 30 can be operated so that measured projection data from the radiation sources 20 and 30 are acquired simultaneously (or essentially / nearly simultaneously, e.g., within about 50 milliseconds from each other) or sequentially (e.g., separated by seconds, minutes, etc.).

[0064] It will be understood that the radiation sources 20, 30 and the detectors 24, 34 may be configured in several ways to provide rotation around the patient during imaging and / or therapeutic scanning. In one embodiment, the movement and exposure of the radiation sources 20, 30 can be synchronized with the longitudinal movement of the patient support 18 to provide continuous helical acquisition or scanning of patient images during the procedure. In addition to the continuous rotation of the radiation sources 20, 30 and the detectors 24, 34 (e.g., continuous rotation of the gantry relative to a constant patient movement velocity), it will be understood that other modifications can be employed without departing from the scope of the disclosed technology. For example, the rotatable gantry 12 and the patient support may be controlled so that the gantry 12 rotates in a “reciprocating” manner (e.g., alternating clockwise and counterclockwise) around the patient supported on the patient support (as opposed to continuous rotation as described above), while the support is controlled to move relative to the rotatable gantry 12 (at a constant or variable speed). In another embodiment, a continuous step-and-shoot circular scan is used, in which longitudinal movement (steps) of the patient support unit 18 is alternated with scanning rotation (shoots) by the rotatable gantry 12 until the desired volume is captured. The multimodal device 10 is capable of volume-based and planar-based imaging acquisition. For example, in various embodiments, the multimodal device 10 may be used to acquire volumetric and / or planar images and perform related processing, including the scattering estimation / correction methods described below.

[0065] Various other types of motion of the radiation source and / or patient support may be utilized to achieve relative motion of the radiation source and / or patient and generate projection data. Discontinuous motion, continuous but variable / non-constant (including linear and non-linear) motion, velocity, and / or trajectory of the radiation source and / or patient support, and combinations thereof may be used, including in combination with various embodiments of the apparatus 10 described above.

[0066] In one embodiment, the rotation speed of the gantry 12, the speed of the patient support unit 18, the beamformer shape, and / or the detector readout value can all be kept constant during image acquisition. In other embodiments, one or more of these variables may change dynamically during image acquisition and / or treatment. The rotation speed of the gantry 12, the speed of the patient support unit 18, the beamformer shape, and / or the detector readout value can be changed to balance different factors, such as image quality, image acquisition time, dose, and workflow.

[0067] In other embodiments, these features can be combined with one or more other image-based activities or procedures, such as patient setup, adaptive therapy monitoring, and treatment planning.

[0068] Afterimages, lag, and ghosting occur. Device 10 and other CT scanning devices typically use detectors such as detector 34, which includes a flat panel. This configuration is shown in Figure 2B.

[0069] During image data acquisition, as the gantry 12 rotates, both the radiation source (e.g., source 30) and the detector (e.g., 34) move along the angular path. Figure 2B shows the angular path 202 for source 36 and the corresponding angular path 204 for detector 34. Source 36 and detector 34 can move synchronously along their respective paths. Data is collected at various positions along the angular path, e.g., positions 204a to 204d.

[0070] At each position 204a to 204d, the detector will spend a period of time collecting the appropriate type of data. The data type may be image data I, which is data relating to the cross-section of X-rays from the X-ray source 36 through the object of interest (not shown). Image data I tends to be diagnostic data relating to a patient. However, in the context of this disclosure, other data may be acquired. Image data I can be reconstructed using tomography to form a 3D representation of the patient. In order to detect the image data I, the source 36 must be powered on and emit X-rays. Alternatively, the data type may be background B data. Background B data is data detected by the detector 34 when the source 36 is powered off. The background B data source may be the residual effect of X-rays on the detector, as will be described in more detail below.

[0071] For many reasons, flat-panel detectors (e.g., detector 36) tend to be made using materials that can trap charges generated by X-ray interactions with the detector. An example of such a material is amorphous silicon. This charge trap can cause a residual or "ghosting" afterglow effect from one image I data acquisition to the next as detector 34 traverses the angular path 204. Ghosting is a spurious signal corresponding to the continued influence of radiation from the previous image I data acquisition during image I data acquisition. This can significantly degrade image quality and hinder quantification. While charge trapping is described herein in relation to flat-panel detectors and detectors made of amorphous silicon, it should be understood that charge trapping is common and can occur in other systems such as silicon photon multipliers (SiPMs). It should be understood that the methods and solutions for ghosting applied herein are not specific to any detector type or material. They can be applied to any detector that exhibits a charge trapping effect.

[0072] Residual signals due to charge trapping may be detected in the current frame from detector exposure in the previous frame. When a large angular range is scanned over a short period of time (e.g., angular path 204 is scanned over several minutes or seconds), ghosting from X-ray exposure at the beginning of the range (e.g., position 204a) can degrade the signal at the later end of the range (e.g., position 204d). This can introduce artifacts into the reconstructed image. Examples of such artifacts include contrast reduction, image heterogeneity, CT number bias (error), and skin line artifacts.

[0073] Ghosting can be particularly severe in IGRT systems (e.g., instrument 10) that include an MVX source (e.g., source 20) because the MVX source generates such a high-energy beam. MeVCT (MVCT) systems can experience cross-contamination, where scattering from the MV beam contaminates cone-beam computed tomography (CBCT) images. This tends to degrade the quality and quantification of CBCT images. Such effects affect CBCT imaging in various ways, including by degrading acquired data and negatively impacting scattering correction using collimator shadow fitting. In radiotherapy (RT) systems using MeV beams, if the keV imaging detector is not adequately shielded from MeV scattering and the keV imaging time is close to the MeV beam on / off cycle, scattering from the MeV can cause significant lag and ghosting in the keV images. If the MeV beam is turned on while keV imaging is acquiring data, scattering from the MeV beam can directly degrade the keV signal acquired while the MeV beam is on. Lag and ghosting can affect more keV data after the MeV beam is turned off.

[0074] When using collimator shadow fitting techniques for scattering correction, lag and ghosting can degrade the collimator shadow data. This can lead to misfitting and misestimation of scattering in the collimator aperture region. Misfitting / estimation can result in inaccurate scattering correction for CBCT imaging.

[0075] In conventional lag correction, images from angles preceding the current imaging angle are weighted by the detector's measured time decay response. These weighted images are then subtracted from the current image. Challenges in this method include the possibility that the conditions during the measurement of a given time decay response may not correspond to the context of patient studies with highly heterogeneous structures and contrasts. These conditions include X-ray energy and the intensity of X-ray irradiation on the detector. Another challenge is that the time interval used for lag correction is empirically determined based on the time decay response. Large time intervals can reduce the accuracy of lag estimation and correction. Yet another challenge arises when the MeV beam is turned on while the CBCT detector is acquiring data (e.g., in an RT scenario). This causes downscattering from the MeV beam, flashing the CBCT detector and resulting in a significant lag effect on data at subsequent acquisition angles. Conventional lag correction methods can be inaccurate.

[0076] Figures 2C to 2F illustrate an exemplary lag introduction problem with a data-driven scattering correction technique for CBCT in the context of this disclosure. Figure 2E shows the target image acquired 1.5 minutes after the image in Figure 2C. Figure 2D shows the profile of the image in Figure 2C. Figure 2E shows the profile of the image in Figure 2D. Note that the coordinate directions on the distance axis are reversed between Figure 2C and Figure 2D. Also, they are reversed between Figure 2E and Figure 2F. This explains why features 206 and 208 appear at the top of Figures 2C and 2E, respectively, but on the left side of Figures 2D and 2F.

[0077] More specifically, the image in Figure 2C was acquired using a collimator (not shown) with a relatively wide aperture. This wide aperture allowed region 206 to be directly irradiated by an X-ray source (e.g., source 30). The irradiation was strong enough to cause charging in the detector (e.g., detector 34). This charging caused an image lag between Figure 2C and Figure 2E. This lag manifests as an increase in detected photon intensity 208 in Figure 2E (profile Figure 2F), even though Figure 2E was acquired more than one minute after Figure 2C. The increase 208 directly corresponds to region 206. This is the residual effect of the over-irradiated region 206 in Figure 2C. Figure 2E was acquired with a slightly smaller collimator aperture. The increase 208 due to the detector lag resulted in a significant underestimation of scattering using the collimator shadow fitting technique.

[0078] As will be discussed below, the algorithms disclosed herein can, among other things, improve or even eliminate these charging effects. One method is to measure the effect of irradiation within region 206 on the detector while the source 36 is off, before obtaining Figure 2E. This is referred to as the “background” B image or data. This data can then be used to subtract or otherwise remove background B from Figure 2E to eliminate the lag effect.

[0079] Improved lag compensation In exemplary variants, one or more background B images are acquired at specific angles of rotation of the gantry (e.g., gantry 12) while the X-ray source is turned off. In the variants, these background B images are used to estimate lag correction accuracy by comparing them to time-decayed response weighted preceding images. Background images may also be used to calibrate the lag correction (e.g., per pixel). This may allow for handling high levels of heterogeneity in patient studies. Background images can also be used to correct MeV scattering. In this case, the background images may be acquired after the MeV beam has been turned off. After the subsequent background images have been acquired, the lag correction may be calibrated to more accurately handle MeV beam-down scattering. Scattering correction can be improved by using collimator shadowing techniques. Other data-driven scattering correction techniques in RT may be adapted to handle lag from MeV down scattering. The variants described herein include improved techniques for lag correction in CBCT imaging using imaging of background B data.

[0080] Background measurement during image acquisition Figure 3A shows the rotation path 202 for the radiation source 30 and the rotation path 204 for the detector 34 within the gantry 12. Figure 3A divides the rotation path 202 into various angular slices (e.g., angular slices 302a to 302d). Note that in this specification, the term “slice,” as in “angular slice,” is used interchangeably with the term “location” or “position” along either of the paths 202 or 204. Understand that a particular location along a path corresponds to a particular angular slice, and vice versa.

[0081] The sizes of the angular slices (e.g., 302a and 302b) are illustrative only. They may include portions of the path 202 that are larger or smaller than those shown in Figure 3A. In general, there may be a correlation between the slice size and the rotation speed of the detector 34 and the source 30 (i.e., the speed at which the source 30 traverses the path 202), which may also depend on other factors such as the size of the gantry 12, the energy of the source 30, and the purpose of the scan. Any suitable rotation speed is possible. Exemplary rotation speeds include 1 rpm, 5 rpm, 7 rpm, and 10 rpm, 20 rpm, and 60 rpm. As shown in Figure 3A, the angular slices do not need to be uniform.

[0082] Each angular slice in Figure 3A is labeled "I" when imaging data is collected in the slice. If background data is collected in the slice, the angular slice is labeled "B" in a circle. While the source 30 / detector pair is positioned to correspond to the image I slice, source 30 supplies imaging radiation. When source 30 / detector 34 is positioned to correspond to the background B slice, source 30 is turned off. To maximize imaging data I, it may be advantageous to have as few background B slices as possible on path 202. Therefore, strategically determining the position of the background B slide for maximum correction effectiveness is important.

[0083] Figure 3A shows exemplary background B positions in image acquisition. Other positions may be used. Figure 3A shows six background B positions, but there may be more or fewer. The exact number may depend on several factors, including the rotation speed of the radiation source, the resolution of the detector, the size of the gantry 12, and the length of the path 202. In one exemplary variant, there is one background B measurement for every 29 angular slices. Background B scans may occur every 7, 15, 30, or 50 slices, or more. Some measurements may have "interleaved" background measurements (i.e., every other background on the path: I / B / I / B / I / B...). In general, the number of background B slices will depend on the angular resolution of the scan. In one variant, 5% of the slices on paths 204 and 202 can be background B. In another embodiment, the background may occupy 1%, 2%, 10%, 15%, 20%, or even 30% of the path and / or rotation. However, it should be understood that there is an inverse correlation between the proportion of paths allocated to imaged I data and background B data. Therefore, there is a trade-off between improved imaging through background B correction and the collection of more imaged I data. For example, the more background B correction occurs during rotation, the less imaged I data is available for imaging. Therefore, in many cases, it is preferable to maximize the imaged I data between background B corrected slices.

[0084] In certain cases, position B may be arbitrary or random, especially when the charging effect and background effect are nearly isotropic. Position B may be evenly spaced around the circumference of path 202. They may be positioned to coincide with, or preceding, a specific time event (e.g., treatment or irradiation of other high-energy X-ray radiation from source 20 (Figure 2A)). One example is positioning background B measurements where there is a large change in X-ray irradiation from one frame to the next. This occurs, for example, when the first frame has an image region A that is completely irradiated by X-rays, and the same region A is shadowed by the X-rays in the next frame. This maximum difference in illuminance between frames results in a lag effect with the highest probability. Background B measurements with this lag will have a significant impact on the image I data and are therefore crucial for correction. These possibilities will be discussed in more detail below.

[0085] Figure 3B illustrates the concept of staggering the angular slice positions of background B within a scan set. This shows a comparison between the background B measurement position in the first angular scan 350 and the background B measurement position in the second angular scan 360. As mentioned above, data acquisition can include multiple scans (e.g., two or more scans of detector 34 along path 204). This is done, for example, to average the results and improve accuracy. In these multiple scan sets, one important concept is to ensure that the background B data acquisition positions do not consistently overlap. This overlap results in the loss of image I data from the angular slice of the target object 304. As a result, this data is missing from the final tomographic image, making it incomplete.

[0086] Figure 3B shows how to stagger the background B positions to prevent this overlap. The first scan 350 has the same background B positions as in Figure 3A. Specifically, in the first scan 350, background B is measured at positions 302e, 302f, 302i, 302j, 302l, and 302n. In the second scan 360, the background is measured at intentionally different positions 302a, 302g, 302h, 302k, 302m, and 302o. The exemplary case in Figure 3B shows background B measurements rearranged on adjacent slices between the first scan 350 and the second scan 360, but this is merely illustrative. Any new configuration of background positions is possible, as long as the overlap between the background B positions and previous scans is reduced or minimized.

[0087] Determination of measurement location for background B Different methods may be used to determine which angle slices measure background B and image I data. One method is to use planning CT images taken before treatment and / or imaging. The “planning CT” images used herein are typically high-quality, high-contrast resolution images that provide fine features and contours of the organ for dose planning. They may be used to set the contrast and other features of the X-ray beam and the resulting image. From the planning CT, projection data in CBCT acquisition may be simulated or predicted. Any suitable algorithm may be developed to determine the distribution of background B locations in path 202 based on the planning CT. The placement of background B locations may be optimized for maximum correction effect and / or minimum image I data loss. The placement may also take into account other considerations besides maximizing the correction effect (e.g., region of interest within subject 304, timing of irradiation from various radiation sources, computation time, power consumption, significantly large fluctuations in signal levels on pixels from one view to the next, or over a portion of a view).

[0088] Another consideration regarding the placement of background B measurements is to avoid measuring background B with respect to the conjugate angle (e.g., in both slices 302c and 302d). The light source 30 illuminates the same portion of object 304 at the conjugate angle (e.g., 302c and 302d, as indicated by the illumination path 306). Therefore, it is advantageous to have imaging I data acquisition for at least one conjugate so that any reconstructed 3D image can contain at least one image I data from the corresponding portion of object 304. If the background B image is acquired in both conjugates, e.g., 302c and 302d, then there will be no image I data for this portion of object 304.

[0089] Other parameters may be modified to optimize or improve background B measurement and correction, or for other purposes. Although Figure 3A appears to show paths 202 and 204 as circular, they can and may have other shapes, such as helical or semi-helical, because the patient or object of interest 304 is typically moving in and out of the page (with respect to Figure 3A) while the scan is being performed. Pathways 202 and 204 may also be elliptical, or may have other shapes. Many of the considerations regarding the placement of background B measurement discussed in the context of the circular path shown in Figure 3A apply to these other shaped paths. Additional considerations may also exist. For example, in the case of a helical path, the pitch of the helix (or other parameters) may be changed to improve background B measurement and / or correction.

[0090] Using simulated X-ray projection data (e.g., simulating the projection of X-rays from source 30 through target 204 and detected by detector 34), it is possible to determine which angular slices would benefit most from acquiring background B for lag correction. This determination can be made, for example, by measuring background B when detector 34 receives the highest radiation level. As mentioned above, background measurements B can also be timed to the irradiation of a specific type of radiation (e.g., MeV therapeutic radiation). Angular slices following irradiation of high-energy X-rays may have a lag or charge effect. Therefore, as described below, it is likely to be advantageous to cluster background B near those angular slices to compensate for the background. These angular slices can also be time-adjusted based on the severity of the temporal changes in the irradiation of target 304. The criterion for the distribution of background B when measuring angular slices around path 202 may also include angles with large changes in detected X-ray intensity. These changes may be due to large contrast changes or orientation changes in the patient (e.g., anterior / posterior i.AP, and left / right i.L / R).

[0091] Other possible sources for input to the algorithms discussed herein include 2D survey images taken at a single angle over a specific length of the patient or subject 304. The 2D survey images may include orthogonal angles (e.g., anterior / posterior (AP), left / right (LR)). They may be used to determine the angle from which to acquire background images for lag correction. The background measurement B positions may be determined using any other suitable images. Suitable image sources include image models of the patient or subject of interest, images taken via other CT / X-ray devices, and / or optical images. Using images and / or prior data, patient-specific or subject-specific lag models can be generated. The patient-specific model may incorporate features of the patient's CT scan, either from the same device used to measure the background or from prior images.

[0092] In addition, or alternatively, any of the algorithms described herein may identify angle groups for background B measurement predetermined by means other than image analysis. Examples include algorithms relating to the use of general patient atlases, patient models (general or phenotypic), phantom models, other mathematical models, and / or physical models. As used herein, “patient atlas” represents a mapping of the patient’s contours observed in X-ray imaging. Patient atlases may be obtained from various sources, including those developed based on large-scale collections of patient images acquired with apparatus performing the methods disclosed herein.

[0093] The algorithm described herein can determine an angle slice for measuring background B during acquisition. Specifically, after a certain angle has been acquired, the algorithm analyzes the already acquired image to determine whether the background image should be acquired.

[0094] The angle of background B measurement within the detected data can be identified within the data by adding tags or other data files in the header. These designations can be used to separate background B data from image I data. Designations can be automatically detected using various algorithms. Appropriate detection algorithms include those that generate a total count for each view. The angle of background B can be significantly smaller than the angle of the adjacent angle when the beam is on. For example, in areas where the X-ray is directly irradiated, or in areas near the boundaries of the patient's body, the signal in the background image when the X-ray is off may be two orders of magnitude lower than the adjacent angle when the X-ray is on.

[0095] Background B data may be acquired during a patient scan for detector lag correction. For example, background B data may be acquired during a predetermined acquisition sequence before the scan, or it may be calculated during the scan (e.g., on the fly). In addition, the algorithm can use the patient's measured background B information to automatically generate a favorable or optimal set of background B measurement locations for that patient. Interpolation may be used in this and other algorithms described herein. The total number of background B data acquisitions may be small compared to the total number of projection views (i.e., views B and I associated with angular slices along path 202 in Figures 3A and 3B). Interpolation, in particular, can facilitate accurate image reconstruction. Interpolation can be performed while background projection images can be used to improve lag correction.

[0096] The angular distribution of background B measurements along path 204 can be intentionally altered over time, particularly for different image scans. In one variant, two or more full rotations of source 30 are used to generate image I data for one 3D reconstruction. The positions of background B measurements in the two rotations can be intentionally made different so that every possible angle along path 304 has at least one image I data scan. In other words, the algorithm can ensure that angles or positions along path 204 assigned to background B measurements in the first rotation (as shown, for example, in Figure 3B) can be assigned differently to image I measurements in the second rotation, and vice versa. It should be understood that the discussion of two scans here is merely illustrative. This technique can be used with any appropriate number of scans. It may also be used in helical scans where the background B data acquisition positions are intentionally changed between scans. In any case, reconstruction can estimate data for these angles from data from adjacent angles (e.g., interpolation) and corresponding angles in subsequent or previous rotations.

[0097] The angular distribution of background B measurements can also be determined based on the raw data during imaging. A model describing the dependence of the angular distribution of background B can be based, for example, on the change in the signal measured at each pixel. The model can then impose a background B measurement if that change increases beyond a certain threshold.

[0098] Detection and correction of image I data using background B Detectors (e.g., detectors 24 and 34, which may be flat-panel detectors), an X-ray tube (not shown), and a system control device (e.g., controller 60) can be synchronized to acquire background data B, during which the X-ray sources (e.g., sources 20 and 30) may be turned off. In an RT scenario, when a MeV beam (e.g., from source 20) is turned on, information about the detected radiation is passed to the CT acquisition control device (e.g., element 60 in Figure 2A) so that background B acquisition can begin after the MeV beam is turned off. Alternatively, an algorithm may detect whether a MeV beam is turned on. In that case, images affected by the MeV beam are tagged. A signal may be generated to trigger the CBCT system to acquire one or more background images. If a MeV beam is turned on during a CBCT scan, predetermined angle groups relative to the background images may be adjusted accordingly. An algorithm can be developed to perform this adjustment.

[0099] One advantageous use of background B measurement is by linear interpolation. For example, the first background B measurement can be performed immediately before acquiring image I. The second background B measurement can be performed immediately after acquiring image I. Then, two background measurements adjacent in time and angle to the measurements in image I can be linearly interpolated to create a composite background error image that possesses the features of both measurements. This composite background image can then be used to correct the data in image I with higher accuracy than using one or more other background images.

[0100] Background B correction allows the use of background B measurements in real time. This can be used to update and improve background correction techniques based on time decay models. For example, time decay modeling lag correction used to correct data can be recalibrated with new information obtained from ongoing or new background B measurements. As with other algorithms described herein, this can be achieved, among other things, using interpolation methods. Background B information from multiple background measurements can be combined to provide correction. This can be particularly advantageous when one of the scans provides background B information about a transient state (e.g., when the X-ray source of therapeutic radiation 20 is switched on or has just been switched off).

[0101] In the deformed form, lag correction may be updated with background information. In lag correction, images from a particular angle are weighted using the measured time decay response of the detector used to detect them. The weighted images can then be subtracted from the current image to correct it. The time decay response is most typically determined based on previous detector measurements. The time interval for which previous images are used in lag correction is most typically determined empirically based on the time decay response. The accuracy of the correction depends on these parameters.

[0102] Any of the background measurements described herein can be used in conjunction with time delay correction. For example, time delay correction can be periodically recalibrated based on background images. For example, it can be recalibrated based on all background images, based on a combination of background images, or based on the average of the background images. Time delay correction may include more than one attenuation term.

[0103] Reconstruction of corrected image I data can be performed using any suitable conventional analytical reconstruction. Typically, such algorithms prioritize or require that image I data be acquired from evenly distributed locations along path 204. Such reconstruction schemes can be modified to adapt to irregular background imaging along path 204 so that background B measurement locations can be constructed and / or optimized. Image reconstruction algorithms can be modified to compensate for data loss at angular locations where background B data is obtained. Compensation can be achieved by any suitable method. One example is to use view-by-view analytical reconstruction and / or iterative reconstruction. Such techniques can minimize the impact of a subset of angles used solely for background B data acquisition.

[0104] Reconstruction may involve data from multiple scans and rotations, as described above. Image reconstruction algorithms may include projection angle sampling. In particular, the azimuth position of the background B image in the current rotation differs from the azimuth position of the background B image in other rotations, especially those occurring before and after the current rotation. In this and other cases, it may be preferable to interleave the background B measurements from rotation to rotation, and possibly to the maximum extent possible. Data redundancy and helical pitch may be used in this consideration, as described above.

[0105] This method allows for modifications to more conventional reconstruction algorithms. For example, conventional algorithms may require images that are evenly spaced (angle-wise). This requirement may be relaxed and modified to accommodate spaced background B images according to any of the schemes, algorithms, or examples described herein. The modification may take into account the non-uniform spacing between background acquisition angles and the non-uniform spacing between image acquisition angles. The modification may take into account the difference between azimuth positions of background images with different rotations of the source 30. In some variations, azimuth positions of one rotation may be interleaved with azimuth positions of another rotation.

[0106] Exemplary implementation of the algorithm Figures 4A and 4B show exemplary algorithm 400, which may perform any of the steps and techniques described herein to correct image lag or ghosting (Figures 2B to 2F). It should be understood that algorithm 400 is general and may incorporate any of the methods or variations of the algorithm described above. Specific steps are shown, but they are not required. Additional steps may be added to algorithm 400 and may still be within the context of this disclosure. Algorithm 440 may be performed using any suitable apparatus, including all the apparatus described herein (e.g., apparatus 10 in Figure 1). It may also be performed using an apparatus having imaging capabilities only. It may also be performed using an apparatus operating in an RT configuration.

[0107] In step 410, algorithm 400 determines the angular path that the X-ray source traverses during imaging. This path can be predetermined by the physical configuration of the imaging device (for example, so that path 202 is for source 30 as shown in Figures 2B and 3). In some variations, the path can be modified by algorithm 400. For example, if the path is helical, the helical pitch can be changed, for example, to improve background B measurements. Any other suitable path determination or modification is possible.

[0108] In step 420, background B measurement locations may be selected along the path identified in step 410. Any suitable background B measurement locations may be selected according to any of the considerations described herein. Often, it may be advantageous to minimize or substantially limit the number of background B measurement locations in order to maximize the image I information acquired. In other embodiments, as described above, it may be advantageous to increase the background B measurements to capture specific temporal phenomena (e.g., switching to or from the X-ray source). As described above, background B acquisition locations may be selected for one or more rotations.

[0109] As described above, the background B position can be selected via an algorithm. The algorithm can make the selection based on any appropriate number of inputs, including the current image and previous images. In this step, any other inputs or considerations may be made to select the background B position. Figure 4A shows step 420, which is performed before data acquisition (step 440), but step 420 is not limited to this order. Step 420 may be performed before or between steps 440. Any suitable algorithm, including machine learning algorithms such as artificial neural networks and decision trees, may perform this step. Other numerical regression techniques may also be used.

[0110] In step 430, the target image acquisition location is selected. In many cases, only measurement locations for background B and image I are possible, so this step is effectively achieved by selecting the background B measurement location in step 420. In other embodiments, there may be locations that do not correspond to either background B or image I measurement. For example, if certain transients that cannot be effectively compensated for through background B subtraction are affecting the X-ray detection signal, it may be advantageous not to acquire data, or to acquire and discard data. These may include transients that occur too frequently for effective background B measurement.

[0111] In step 440, the X-ray source (e.g., source 30) is moved along the path selected in step 410 to acquire an image. The acquired images include both background B data and image I data. The location where this data is acquired will follow step 420.

[0112] In step 450, errors in the image data I (e.g., lag or ghosting effects as shown in Figures 2B to 2F) may be corrected using background data B. This correction can be carried out by any of the methods described herein. One example is simply subtracting the background B data image from image I. Other examples include interpolation techniques, for example, interpolating two or more background images in combination with subtraction, or interpolating them without subtraction. Still other techniques may include selectively modifying pixels in the image (e.g., by convolution) based on problems identified through the background images. Any suitable algorithm may perform this step, or it may be performed manually. Suitable algorithms include machine learning algorithms such as artificial neural networks and decision trees. Other numerical regression techniques may also be used.

[0113] In step 460, a corrected image is generated using the correction results from step 450. This image may be individual 2D images of a particular scan. This image is a 3D image containing many such 2D images. It may also be a tomographic reconstruction. The corrected image can be used for any appropriate purpose, including patient diagnosis and treatment. The corrected image may also be used to aid in the generation of background B measurement locations and / or to inform any other aspect of the algorithm 400 described above.

[0114] Figure 4B shows details of the image acquisition step 440 in algorithm 400. Acquisition step 440 may be repeated over one or more movements of the X-ray source. Step 440 is generally performed until a complete set of imaging I data is acquired. Whether the set is "complete" is relevant to the task at hand. This set may contain one or more images I corresponding to one or more rotations. Typically, the rotations traverse the entire path identified in step 410. However, it should be understood that this is not required for all implementations of step 440.

[0115] In step 441, the radiation source (e.g., source 30) moves along an angular path to a specific position. In step 442, algorithm 400 determines whether the current position corresponds to the acquisition of background B or image I. As described above, other options may be available (e.g., some positions may be flagged so that data acquisition does not occur at other positions in order to allow for transient disappearance). The determination of the type of data to be acquired can be based on the use of any of the algorithms or considerations described above. Once the data type is selected, the algorithm sets the operating state of the X-ray source in step 443 according to the data type selection. If the data to be acquired is background B data, the source is switched off, or remains powered off if it is already off. If the data to be acquired is image I data, the source is powered on if it is not powered on, or remains powered on. Once the source operating state is set, in step 444, the appropriate type of data is acquired and stored. Algorithm 400 may wait to acquire data until the transient caused by switching the source operating state has disappeared. Once data has been collected, algorithm 400 queries in step 445 whether the source has reached the end of the path. If so, algorithm 400 proceeds to step 450 (Figure 4A) to remove errors from the collected data. Otherwise, algorithm 400 returns to step 441 and moves the source to a new position. Algorithm 400 repeats this process until the source reaches the end of the path specified in step 410.

[0116] When the above apparatus and method are used in the projection domain, they can be applied to each projection view, and each projection view is a planar image. Various embodiments can utilize different scan geometries, detector positions (including offset detectors), and / or beam shaping window shapes.

[0117] As described above, embodiments of the disclosed technology can be used in radiotherapy devices and methods that utilize multimodal radiation sources, including integrated low-energy (e.g., keV) and high-energy (e.g., MeV) sources, for use with or as part of IGRT. According to one embodiment, the imaging acquisition method includes, for example, a helical source trajectory (e.g., continuous source rotation around a central axis with longitudinal movement of a patient support through a gantry bore) or a circular scan, or can be used in other ways, along with high-speed slip ring rotation, to provide keVCT imaging on a radiotherapy delivery platform.

[0118] While the disclosed technology is shown and described in relation to specific aspects, one embodiment, or more embodiments, it will be apparent to those skilled in the art that equivalent modifications and changes will occur in reading and understanding this specification and the accompanying drawings. In particular, with respect to the various functions performed by the elements described above (components, assemblies, devices, members, compositions, etc.), the terms used to describe such elements (including references to “means”) are intended, unless otherwise indicated, to correspond to any element that performs a particular function of the described element (i.e., is functionally equivalent) to a disclosed structure that performs that function in an exemplary aspect, embodiment, or embodiment of the disclosed technology illustrated herein, even if it is not structurally equivalent to such structure. In addition, while certain features of the disclosed technology may be described above in relation to only one or more of the illustrated aspects or embodiments, such features may be combined with one or more other features of other embodiments so as to be desirable and advantageous for any given or particular application.

[0119] The embodiments described herein relate to the systems and methods described above, but these embodiments are intended to be illustrative and not to limit the applicability of these embodiments to the descriptions herein. The present invention is illustrated by the description of its embodiments, and these embodiments are described in some detail, but it is not the applicant's intention to limit the appended claims to such detail or to limit them in any way. Further advantages and modifications will be readily apparent to those skilled in the art. Accordingly, the present invention, in its broader embodiments, is not limited to the specific details illustrated and described, representative apparatus and methods, and exemplary embodiments. Accordingly, deviations from such details can be made without departing from the spirit or scope of the applicant's general inventive concept.

Claims

1. It is an imaging system, The first X-ray source, To generate a first X-ray photon in a first energy range suitable for imaging, Projecting the first X-ray photon onto a region designated for imaging, The first X-ray source configured to perform the following: A rotatable gantry configured to rotate the first X-ray source so that the first X-ray source traverses an angular path, A data processor, Determining the set of background acquisition angles at least partially by determining the minimum number of measurement positions for background data and maximizing the number of image acquisition angles between background acquisition angles, To collect first data relating to the transmission of the first X-ray photon through the region designated for imaging, using the set of image acquisition angles along the angular path, Collecting background data in a set of background acquisition angles along the aforementioned angular path, wherein the system acquires image slices of two or more image acquisition angles in the designated region for imaging between adjacent background acquisition angles. Using the aforementioned background data, errors in the first data are removed, Based on the removal of the error in the first data, a corrected image is generated. The data processor has an analysis section configured to perform the following: An imaging system equipped with the following features.

2. Using the aforementioned background data, errors can be removed from the first data. Using the measured time decay response of the X-ray detector, the background image from a specific angle is weighted, Subtracting the weighted background image from the first data, The system according to claim 1, including the following:

3. The system according to claim 2, wherein the analysis unit recalibrates the time decay response based on the background data.

4. The system according to claim 3, wherein the recalibration of the time decay response includes two or more decay terms based on two or more background images.

5. The system according to any one of claims 1 to 4, wherein removing errors in the first data using the background data includes interpolating the background image.

6. It is an X-ray source, To generate X-ray photons in an energy range suitable for imaging, Projecting the aforementioned X-ray photons onto a region designated for imaging, The X-ray source configured to perform the following: A rotatable gantry configured to rotate the X-ray source so that the X-ray source traverses an angular path, A data processor, Collecting background data at a first angle along the aforementioned angular path, To collect image data relating to the transmission of the X-ray photons through the region designated for imaging over an angular range along the aforementioned angular path, The collection of background data at a second angle along the aforementioned angular path, wherein the second angle is positioned such that the angular range lies between the first angle and the second angle. A corrected image is generated by removing errors in the image data using interpolation of the background data collected at the first and second angles. The data processor has an analysis section configured to perform the following: An imaging system equipped with the following features.

7. The system according to claim 6, wherein the analysis portion determines the first angle, the second angle, and the angle range at least partially by determining the minimum number of measurement locations for the background data for accurate background estimation.

8. The system according to claim 7, wherein the analysis portion determines the first angle, the second angle, and the angle range at least partially by maximizing at least one of the total number of image slices of the region designated for imaging and the number of image slices of the region designated for imaging taken between the first angle and the second angle, respectively.

9. Removing errors from the aforementioned image data is Weighting the background image using the measured time decay response of the X-ray detector, Subtracting the weighted background image from the image data, The system according to any one of claims 6 to 8, including the system described in any one of claims 6 to 8.

10. The system according to claim 9, wherein the analysis unit recalibrates the time decay response based on the background data.

11. The system according to any one of claims 1 to 10, comprising a cone-beam computed tomography (CBCT) system.

12. The second X-ray source is, The method involves generating a second X-ray photon in a second energy range that has a different energy from the first energy range, Projecting the aforementioned second X-ray photon onto a region designated for imaging, The second X-ray source, configured to perform the following, The aforementioned analysis section is Combining the data obtained from the first X-ray photon and the second X-ray photon, Interleaving the irradiation from the first X-ray source and the second X-ray source, and To operate the first X-ray source and the second X-ray source simultaneously, A configuration which is configured to perform at least one of the following, according to any one of claims 1 to 5 The system described.

13. The system according to any one of claims 1 to 5, wherein the set of background collection angles is evenly distributed along the angle path.

14. The system according to claim 1, wherein the set of background collection angles is distributed at predetermined angles along the angle path.

15. The system according to any one of claims 1 to 5, wherein the set of background acquisition angles is determined via at least one of the following: analysis of planned computed tomography (CT) images, analysis of previously available CT images, analysis of CT2D survey images, analysis of orthogonal angle images, analysis of patient atlases, and analysis of data being acquired.

16. The system according to any one of claims 1 to 5, wherein the set of background acquisition angles is determined using at least one of the patient's geometric shape, acquisition protocol, and projection angle relative to the patient's orientation.

17. The system according to any one of claims 1 to 5, wherein the analysis portion determines the set of background collection angles via an algorithm.

18. The system according to claim 17, wherein the algorithm determines the set of background acquisition angles based on improving the accuracy of CBCT reconstruction.

19. The system according to any one of claims 17 and 18, wherein the algorithm excludes conjugate angles from the set of background acquisition angles.

20. The system according to any one of claims 17 to 19, wherein the algorithm selects a background acquisition angle at least in part based on the use of MeV radiation.

21. The system according to any one of claims 17 to 20, wherein the algorithm selects a background acquisition angle to improve at least one of lag correction and data-driven scattering correction.

22. Using the aforementioned background data, the error in the first data can be removed. Collecting background data at a first angle along the aforementioned angular path, The first data is collected within the angular range along the aforementioned angular path, Collecting background data at a second angle along the aforementioned angular path, wherein the second angle is positioned such that the angular range lies between the first angle and the second angle. The aforementioned collection, By interpolating the background data collected at the first and third angles, To generate an error image, Using the aforementioned error image, the errors in the first data are removed, The system according to any one of claims 1 to 5, including the system described in any one of claims 1 to 5.

23. The system according to claim 17, wherein the algorithm can generate a first set of background acquisition angles for a first scan and a second set of background acquisition angles for a second scan, wherein the first set and the second set are different.

24. The system according to claim 23, wherein the first set and the second set do not overlap.

25. The system according to claim 17, wherein the angular path is helical.

26. The system according to claim 25, wherein the algorithm determines the background collection angle based at least in part on the pitch size of the helix.

27. The system according to claim 17, wherein the algorithm determines the background acquisition angle at least in part on the basis of improving the accuracy of the reconstructed 3D image using the first data.

28. The system according to claim 17, wherein the image reconstruction selects a different set of background acquisition angles for different rotations of the first X-ray source.

29. The system according to any one of claims 1 to 5, wherein the analysis unit synchronizes the X-ray control, the flat panel detector readout, and the CT scan control so that when the power to the first X-ray source is turned off, the readout of the flat panel detector provides background data.

30. The system according to claim 29, wherein the movement of the first X-ray source along the angular path is not interrupted while the power to the X-ray source is turned off.

31. The system according to any one of claims 1 to 5, wherein the analysis portion is further configured to generate a reconstructed 3D image by modifying the reconstruction algorithm based on the background acquisition angle.

32. The system according to claim 31, wherein the modification takes into account the non-uniform intervals between background acquisition angles and the non-uniform intervals between image acquisition angles.

33. The system according to claim 31, wherein the modification takes into account the difference between the azimuth position of the background image during the first rotation of the first X-ray source and the azimuth position of the background image during the second rotation of the first X-ray source.

34. The system according to claim 33, wherein the azimuth position of the first rotation is interleaved with the azimuth position of the second rotation.

35. A method for operating an imaging system, To generate a first X-ray photon in a first energy range suitable for imaging, Projecting the first X-ray photon onto a region designated for imaging, Rotating the first X-ray source so that it traverses the angular path, Determining the set of background acquisition angles at least partially by determining the minimum number of measurement positions for background data and maximizing the number of image acquisition angles between background acquisition angles, To collect first data relating to the transmission of the first X-ray photon through the region designated for imaging, using the set of image acquisition angles along the angular path, Collecting background data in a set of background acquisition angles along the aforementioned angular path, wherein the system acquires image slices of two or more image acquisition angles in the designated region for imaging between adjacent background acquisition angles. Using the aforementioned background data, errors in the first data are removed, Based on the removal of the error in the first data, a corrected image is generated. Methods that include...

36. The first X-ray source, To generate a first X-ray photon in a first energy range suitable for imaging, Projecting the first X-ray photon onto a region designated for imaging, The first X-ray source configured to perform the following: A rotatable gantry configured to rotate the first X-ray source so that the first X-ray source traverses an angular path, A data processor, Determining the set of background acquisition angles at least partially by determining the minimum number of measurement positions for background data and maximizing the number of image acquisition angles between background acquisition angles, To collect first data relating to the transmission of the first X-ray photon through the region designated for imaging, using the set of image acquisition angles along the angular path, The background data is collected using the set of background collection angles along the aforementioned angular path, Using the aforementioned background data, errors in the first data are removed, Based on the removal of the error in the first data, a corrected image is generated. The data processor has an analysis section configured to perform the following: An imaging system equipped with the following features.

Images