System and method for CT detector calibration using a wire phantom

Abstract

To provide systems and methods for calibrating a computed tomography (CT) system.SOLUTION: In one example, a method for a computed tomography (CT) system comprises, during a calibration of the CT system, measuring a position of a detector element of a detector array of the CT system using a wire of a wire phantom coupled to a table of the CT system, during a rotational scan performed using the CT system; during a subsequent scan performed on a subject using the CT system, applying the measured position of the detector element rather than a design target position of the detector element during reconstruction of an image from projection data acquired via the CT system (508); and displaying the reconstructed image on a display device of the CT system (510).SELECTED DRAWING: Figure 5A

Description

[Technical Field]

[0001] FIELD OF THE INVENTION Embodiments of the subject matter disclosed herein relate to computed tomography (CT) imaging systems, and in particular to the calibration of CT imaging systems using a wire phantom. [Background technology]

[0002] In a computed tomography (CT) imaging system, a beam of electrons generated by a cathode travels toward a target in an x-ray tube. The electrons strike the target, producing a fan- or cone-shaped x-ray beam that is directed toward an object being examined (such as a patient). After the x-rays are attenuated by the object, they strike an array of radiation detectors. Each detector element generates an electrical signal, which is used to reconstruct an image of the object. Each electrical signal corresponds to a voxel / pixel in the image.

[0003] Image quality in terms of resolution, contrast-to-noise ratio, and other factors depends on the accuracy of the alignment of each detector element within the detector array. Misalignment of detector elements can increase image artifacts and / or degrade image quality. Misalignment can be measured and corrected using mechanical alignment techniques. Mechanical alignment techniques can employ several types of measurement methods, for example, using contact-type and / or laser-based positioning tools. Based on the measurement data, the detector elements can be realigned using a set of mechanical tools. However, current alignment performance is not suitable for newer versions of CT systems, which include smaller detector elements, requiring alignment accuracy finer than 10 μm. Summary of the Invention

[0004] The present disclosure addresses at least one or more of the above-identified problems with a method for a computed tomography (CT) system. The method includes measuring positions of detector elements of a detector array of the CT system using wires of a wire phantom coupled to a table of the CT system during calibration of the CT system during a rotational scan performed using the CT system; applying the measured positions of the detector elements, rather than design target positions of the detector elements, during image reconstruction from projection data acquired by the CT system during a subsequent scan performed on a subject using the CT system; and displaying the reconstructed image on a display device of the CT system. The measured positions of the detectors can be calculated for each detector element of each detector array of the CT system and stored as calibration vectors in the memory of the CT system. The calibration vectors can be used to mechanically adjust large detector misalignments up to the precision limit of the mechanical adjustment (e.g., greater than 20 μm). The calibration vectors can also be applied to the output of the detector elements during image reconstruction to correct for detector misalignments smaller than the precision limit. For example, some detector elements may become misaligned before or during installation of the CT system or after replacing an inoperable portion of the detector array.

[0005] Applying the measured detector positions, rather than the design target positions, to subsequent scans of the subject can improve the quality of the resulting reconstructed image. An advantage of the above-described method is that, unlike cases where only limited mechanical alignment techniques can be relied upon to adjust the alignment of detector elements, measured detector position values ​​are calculated using projection geometry based on measuring the position of each detector relative to a known reference point (e.g., a wire) at different view angles and applied to image reconstruction. As a result, the method can be applied to detector arrays containing smaller detector elements than can be aligned using mechanical alignment techniques.

[0006] Furthermore, the elimination of other precision mechanical adjustments and associated calibration vectors and processes reduces CT system calibration time, resources, and costs, thereby increasing CT system efficiency. This improves CT system functionality, reduces downtime between calibrations and exams performed using the CT system, and allows for faster, more accurate processing. Furthermore, application of the calibration methods described herein enables CT systems to produce higher quality images, resulting in more successful diagnoses and more desirable patient outcomes.

[0007] These and other advantages and features of the present specification will be readily apparent from the following detailed description alone or from the drawings and the following detailed description. It should be understood that the above summary is provided to select concepts further described in the detailed description in a simplified form. It is not intended to identify key features or essential features of the claimed subject matter, the scope of which is defined solely by the claims. Moreover, the claimed subject matter is not limited to implementations that solve the above-described shortcomings or shortcomings noted in any part of this disclosure. [Brief explanation of the drawings]

[0008] The various aspects of the present disclosure will be better understood by reading the following detailed description and by reference to the drawings, in which: [Figure 1] FIG. 1 is a perspective view of a CT imaging system in accordance with one or more embodiments of the present disclosure. [Figure 2] FIG. 1 is a block schematic diagram of an exemplary CT imaging system in accordance with one or more embodiments of the present disclosure. [Figure 3A] 1 is a wire phantom for calibrating a CT imaging system, in accordance with one or more embodiments of the present disclosure. [Figure 3B]3B illustrates the alignment of the wire phantom of FIG. 3A with respect to the table of a CT system in accordance with one or more embodiments of the present disclosure. [Figure 3C] 3B illustrates the alignment of the wire phantom of FIG. 3A with respect to the detector array of a CT system in accordance with one or more embodiments of the present disclosure. [Figure 4] 1 is a schematic diagram of an exemplary detector array of a CT system in accordance with one or more embodiments of the present disclosure. [Figure 5A] 1 is a flowchart illustrating an exemplary method for improving the accuracy of CT images using calibration data, in accordance with one or more embodiments of the present disclosure. [Figure 5B] 1 is a flowchart illustrating an exemplary method for calibrating a CT system using a wire phantom, in accordance with one or more embodiments of the present disclosure. [Figure 6] 1 is a schematic diagram illustrating an exemplary alignment of a wire phantom with respect to an X-ray source and an X-ray detector array, in accordance with one or more embodiments of the present disclosure. [Figure 7] 10 is a graph of a sinogram generated using a wire phantom in accordance with one or more embodiments of the present disclosure. [Figure 8] FIG. 10 illustrates how a parabolic function can be fitted to the output of a detector array to determine the position of a detector element in accordance with one or more embodiments of the present disclosure. [Figure 9] 10 illustrates an exemplary projection geometry used to calculate the positions of detector elements of a detector array relative to a wire phantom, in accordance with one or more embodiments of the present disclosure.

[0009] The drawings illustrate certain aspects of the described systems and methods. Together with the following description, the drawings illustrate and explain the structures, methods, and principles described herein. In the drawings, the dimensions of components may be exaggerated or otherwise modified for clarity. Well-known structures, materials, or operations have not been shown or described in detail to avoid obscuring aspects of the described components, systems, and methods. DETAILED DESCRIPTION OF THE INVENTION

[0010] SUMMARY OF THE INVENTION

[0003] The present specification and embodiments of the subject matter disclosed herein relate to methods and systems for improving the quality of images acquired through a computed tomography (CT) system. Generally, in a computed tomography (CT) imaging system, an X-ray source emits a fan beam or a cone beam toward an object (such as a patient). The X-rays emitted by the X-ray source are attenuated to various levels by the object before being detected by radiation detector elements arranged in one or more detector arrays. In a CT system, the X-ray source and detector array typically rotate around a gantry within an imaging plane and around the patient, and an image is generated from projection data in multiple views at different view angles. For example, 1000 views can be generated by the CT system for one rotation of the X-ray source. After being attenuated by the patient, the beam impinges on the array of radiation detector elements. The detector array typically includes a collimator for collimating the x-ray beam received at the detector, a scintillator disposed adjacent to the collimator for converting x-rays into light energy, and a photodiode for receiving the light energy from the adjacent scintillator and generating an electrical signal. The intensity of the attenuated beam radiation received at the detector array typically depends on the attenuation of the x-ray beam by the patient. Each detector element of the detector array generates a separate electrical signal indicative of the attenuated beam received by the detector element. The electrical signals are transmitted to and analyzed by a data processing system. The data processing system processes the electrical signals to generate an image.

[0011] The detector array can be curved so that detector elements disposed on the detector array are aligned toward the x-ray source, thereby enabling more accurate and efficient detection of x-rays emitted by the x-ray source. The detector elements can be arranged in a cluster format. For example, the detector array can include multiple detector modules, each of which can include multiple groups of detector elements called chiclets, each of which can include multiple detector elements.

[0012] The quality of an image reconstructed from projection data collected from a detector element may depend on the alignment of each detector element with the x-ray source. Some detector elements may not be accurately positioned during assembly due to limited performance of mechanical alignment tools, and may become misaligned over time or through use of the CT system. For example, if a detector module in a detector array is damaged, the detector module is replaced with a new detector module. The new detector module may not be accurately positioned due to the lack of high-precision alignment tools at the customer site. First, the detector module may not be accurately assembled within the detector array, the chiclets may be misaligned within the detector module, and / or the detector elements may be misaligned within the chiclets.

[0013] Misalignment of some detector elements can increase noise artifacts in the image and / or degrade image quality. Correcting misalignment can improve image quality. To correct misalignment, CT systems can be periodically calibrated, and the positions of detector modules, chiclets, or detector elements can be mechanically adjusted. However, the alignment techniques currently used to measure and correct misalignment are not suitable for newer versions of CT systems, which have smaller detector elements and smaller pixel sizes.

[0014] To address the problem of calibrating CT systems containing small detector elements and correcting for detector element misalignment, a method is proposed herein for directly measuring detector element positions. The method generates detector positions from projection data collected using a wire phantom without relying on mechanical alignment techniques. Calibration vectors can be generated from the measured detector positions, and these calibration vectors can be applied to projection data acquired in subsequent scans during image reconstruction to correct for misaligned detector positions. In some embodiments, rather than correcting for misaligned detector positions, the measured detector positions can be applied directly during image reconstruction instead of the designed target detector positions.

[0015] An example of a CT system is shown in FIGS. 1 and 2. FIG. 3A shows an exemplary wire phantom that can be used to calibrate a CT system. The wire phantom can be mounted on the table of the CT system as shown in FIG. 3B. FIG. 3C shows the position of the wires of the wire phantom relative to the x-ray source and detector array of the CT system during the calibration process described herein. The detector array can include multiple detector elements organized into modules, as shown in FIG. 4. FIG. 5B shows how a CT system can be calibrated to correct for detector element misalignment, and FIG. 5A shows how measured detector positions obtained during the calibration process can be applied during image reconstruction to improve the quality of the resulting image. FIG. 6 shows how the wire phantom is positioned within the imaging plane of the CT system during calibration. During calibration, projection data can be collected at the detector elements as the gantry holding the detector array and x-ray source rotates around the wire phantom. With this rotation, the projection data collected during calibration can be displayed graphically as a sinogram, as shown in Figure 7, which can show which detector elements detect the attenuation of x-rays by the wire at a given view angle or frame. The attenuation can be greatest at the detector element indicated by a parabolic function fitted to the detector output data, as shown in Figure 8. The position of the detector element relative to the wire can be measured according to the projection geometry shown in Figure 9.

[0016] 3A, 3B, 4, and 6 illustrate exemplary configurations of the relative positions of various elements. When elements are shown in direct contact with or directly coupled to one another, in at least one example, the elements can be said to be in direct contact with or directly coupled to one another, respectively. Similarly, elements shown contiguous or adjacent to one another can be said to be contiguous or adjacent to one another, respectively, in at least one example. As an example, elements that are in surface contact with one another can be said to be in surface contact with one another. As another example, elements that are spaced apart from one another can be said to be spaced apart from one another, in at least one example, if there is a space between the elements but not other elements. As yet another example, elements that are shown above and below one another, opposite one another, or left and right from one another can be said to be above and below, opposite one another, or left and right. Furthermore, as shown in the figures, in at least one example, the topmost element or the highest point of an element can be said to be the “top” of the element, and the bottommost element or the lowest point of an element can be said to be the “bottom” of the element. As used herein, top / bottom, upper / lower, and above / below may refer to elements relative to the vertical axis of the figure and may be used to describe the relative positions of elements in the figure. Thus, an element shown on top of another element may, in one example, be positioned directly above the other element. As yet another example, the shapes of elements shown in the figures may be referred to as having those shapes (e.g., circular, rectilinear, flat, curved, rounded, chamfered, angled, etc.). Furthermore, elements shown to intersect with one another may, in at least one example, be referred to as intersecting elements or intersecting one another. Furthermore, elements shown within or outside of other elements may, in one example, be referred to as elements shown within or outside of other elements.

[0017] FIG. 1 illustrates an exemplary CT system 100 configured to perform CT imaging. In particular, the CT system 100 is configured to image an object 112 (e.g., a patient), an inanimate object, one or more manufactured parts, and / or a foreign object (e.g., an implant, a stent, and / or a contrast agent present in the body). In one embodiment, the CT system 100 includes a gantry 102, which may further include at least one X-ray source 104 that emits an X-ray radiation beam 106 (see FIG. 2) used to image the object 112 residing on a table 114. Specifically, the X-ray source 104 is configured to emit the X-ray radiation beam 106 toward a detector array 108 positioned on the opposite side of the gantry 102. While only a single X-ray source 104 is shown in FIG. 1, in certain exemplary embodiments, multiple X-ray sources and multiple detectors may be used to emit multiple X-ray radiation beams and acquire projection data at different energy levels corresponding to the patient. In some embodiments, the X-ray source 104 is capable of performing rapid peak kilovoltage (kVp) switching. In the embodiments described herein, the X-ray detector employed is a photon-counting detector that is capable of distinguishing between X-ray photons of different energies.

[0018] In certain embodiments, the CT system 100 further includes an image processing unit 110 that reconstructs an image of a target volume of the object 112 using an iterative image reconstruction method or an analytical image reconstruction method. For example, the image processing unit 110 may reconstruct an image of the target volume of the patient using an analytical image reconstruction method (such as filtered back projection (FBP)). As another example, the image processing unit 110 may reconstruct an image of the target volume of the object 112 using an iterative image reconstruction method such as advanced statistical iterative reconstruction (ASIR), conjugate gradient (CG), maximum likelihood expectation maximization (MLEM), or model-based iterative reconstruction (MBIR). As described further herein, in some examples, the image processing unit 110 may also use analytical image reconstruction methods (such as FBP) in addition to iterative image reconstruction methods.

[0019] In some CT imaging system configurations, an x-ray source emits a cone-shaped beam of x-ray radiation that is collimated to lie within the XYZ plane of a Cartesian coordinate system and is commonly referred to as the "image plane." The x-ray radiation beam passes through the object being imaged (such as a patient or subject). After being attenuated by the object, the x-ray radiation beam impinges on an array of detector elements. The intensity of the attenuated x-ray radiation beam received at the detector array depends on the attenuation of the x-ray radiation beam by the object. Each detector element in the array produces a separate electrical signal that is a measurement of the x-ray beam attenuation at the detector location. The attenuation measurements from all detector elements are acquired separately to create a transmission profile.

[0020] In some CT systems, the x-ray source and detector array are rotated by a gantry in an imaging plane around the object to be imaged so that the angle at which the x-ray beam intersects the object is constantly changing. A group of x-ray radiation attenuation measurements (e.g., projection data) obtained from the x-ray detector array at one gantry angle is called a "view." A "scan" of the object includes a set of views made at different gantry angles, or view angles, during one revolution of the x-ray source and detector.

[0021] FIG. 2 illustrates an exemplary imaging system 200 similar to the CT system 100 of FIG. 1. In aspects of the present disclosure, the imaging system 200 is configured to image a subject 204 (e.g., subject 112 of FIG. 1). In one embodiment, the imaging system 200 includes a detector array 108 (see FIG. 1). The detector array 108 further includes a plurality of detector elements 202. The plurality of detector elements 202 sense an x-ray radiation beam 106 (see FIG. 2) passing through the subject 204 (e.g., a patient) to acquire corresponding projection data. In some embodiments, the detector array 108 is fabricated in a multi-slice configuration including multiple rows of cells or detector elements 202. In such a configuration, one or more additional rows of detector elements 202 are arranged in a parallel configuration to acquire projection data. An exemplary detector array configuration is described in further detail below with reference to FIG. 4.

[0022] In certain embodiments, imaging system 200 is configured to move through different angular positions around object 204 to acquire desired projection data. Thus, gantry 102 and gantry-mounted components may be configured to rotate about center of rotation 206 to acquire projection data at different energy levels, for example. Alternatively, in embodiments in which the projection angle changes relative to object 204 as a function of time, the mounted components may be configured to move along a general curve rather than along a circle.

[0023] As the X-ray source 104 and detector array 108 rotate, the detector array 108 collects data of the attenuated X-ray beam. The data collected by the detector array 108 is pre-processed and calibrated to adjust the data to represent line integrals of the attenuation coefficients of the scanned object 204. The processed data is commonly referred to as a projection. In some embodiments, individual detectors or detector elements 202 of the detector array 108 may include photon-counting detectors that register individual photon interactions into one or more energy bins.

[0024] The acquired set of projection data can be used for reference material differentiation (BMD). During BMD, the measured projections are converted into a set of material density projections. The material density projections can be reconstructed to form a pair of material density maps or material density images and / or contrast maps for each reference material, such as bone, soft tissue, etc. These density maps or density images can be correlated in order to form a 3D volumetric image of the reference material (e.g., bone, soft tissue, and / or contrast agent) in the imaging volume.

[0025] Once reconstructed, the reference material images produced by the imaging system 200 reveal internal features of the subject 204 that are represented by the densities of the two reference materials. The density images can be displayed to show these features. In a traditional approach to diagnosing a medical condition (such as a disease state), or more generally, diagnosing a medical event, a radiologist or physician would review a hard copy or displayed density image to identify features of interest. Such features might include lesions, the size and shape of particular anatomical structures or organs, and other features that would be identifiable in the image based on the skill and knowledge of the individual practitioner.

[0026] In one embodiment, imaging system 200 includes a control mechanism 208 that controls the movement of components, such as the rotation of gantry 102 and the operation of x-ray source 104. In certain embodiments, control mechanism 208 further includes an x-ray controller 210 configured to provide power and timing signals to x-ray source 104. Furthermore, control mechanism 208 includes a gantry motor controller 212 configured to control the rotational speed and / or position of gantry 102 based on imaging requirements.

[0027] In certain embodiments, the control mechanism 208 further includes a data acquisition system (DAS) 214 configured to sample analog data received from the detector elements 202 and convert the analog data to a digital signal for subsequent processing. The DAS 214 may be configured to selectively aggregate analog data from a subset of the detector elements 202 into a so-called macro-detector, as described further herein. The data sampled and digitized by the DAS 214 is transmitted to a computer or computing device 216. Note that the computing device 216, in at least one embodiment, may be the same as or similar to the image processing unit 110. In one embodiment, the computing device 216 stores the data in a storage device or mass storage device 218. The storage device 218 may be, for example, any type of non-transitory memory, and may include a hard disk drive, a floppy disk drive, a compact disk read / write (CD-R / W) drive, a digital versatile disk (DVD) drive, a flash drive, and / or a solid-state storage drive.

[0028] Additionally, the computing device 216 provides commands and parameters to one or more of the DAS 214, the X-ray controller 210, and the gantry motor controller 212 to control system operation (e.g., data acquisition and / or data processing). In certain embodiments, the computing device 216 controls system operation based on operator input. The computing device 216 accepts operator input, including, for example, commands and / or scanning parameters, via an operator console 220 operably coupled to the computing device 216. The operator console 220 may include a keyboard (not shown) or a touch screen to enable an operator to specify commands and / or scanning parameters.

[0029] 2, more than one operator console 220 may be coupled to imaging system 200, for example, to input or output system parameters, request exams, present data, and / or view images. Additionally, in certain embodiments, imaging system 200 may be coupled through one or more configurable wired and / or wireless networks (e.g., the Internet and / or virtual private networks, wireless telephone networks, wireless local area networks, wired local area networks, wireless wide area networks, wired wide area networks, etc.) to multiple displays, printers, workstations, and / or similar devices located locally or remotely within a facility or hospital or at entirely different locations.

[0030] In one embodiment, for example, imaging system 200 includes or is coupled to a Picture Archiving and Communication System (PACS) 224. In an exemplary embodiment, PACS 224 is further coupled to a remote system (radiology information system, hospital information system) and / or an internal or external network (not shown) to allow an operator at another location to provide commands and parameters and / or access image data.

[0031] The computing device 216 uses operator-supplied and / or system-defined instructions and parameters to operate the table motor controller 226. The table motor controller 226 can control the table 114, which can be a motorized table. Specifically, the table motor controller 226 can move the table 114 so that the subject 204 is properly positioned in the gantry 102 to acquire projection data corresponding to a target volume of the subject 204.

[0032] As previously mentioned, DAS 214 samples and digitizes the projection data acquired by detector elements 202. Image reconstructor 230 then performs high-speed reconstruction using the sampled and digitized x-ray data. While FIG. 2 illustrates image reconstructor 230 as a separate entity, in certain exemplary embodiments, image reconstructor 230 may form part of computing device 216. Alternatively, image reconstructor 230 may not be present in imaging system 200; instead, computing device 216 may perform one or more functions of image reconstructor 230. Furthermore, image reconstructor 230 may be located locally or remotely and may be operably connected to imaging system 200 using a wired or wireless network. In particular, in one exemplary embodiment, computing resources in a “cloud” network cluster may be used for image reconstructor 230.

[0033] In one embodiment, image reconstructor 230 stores the reconstructed image in storage device 218. Alternatively, image reconstructor 230 may transmit the reconstructed image to computing device 216 for generating patient information useful for diagnosis and evaluation. In some embodiments, computing device 216 may transmit the reconstructed image and / or patient information to a display or presentation device 232 communicatively coupled to computing device 216 and / or image reconstructor 230. In some exemplary embodiments, the reconstructed image may be transmitted from computing device 216 or image reconstructor 230 to storage device 218 for short-term or long-term storage.

[0034] The detector array 108 may include multiple detector modules, each of which may include multiple cycles, where a cycle is a set of individual sensors or detector elements.

[0035] 4, an exemplary detector array 400 is shown. It should be understood that the detector array 400 may be configured in different sizes and / or shapes (such as square, rectangular, circular, or another shape). The actual field of view (FOV) of the detector array 400 may be directly proportional to the size and shape of the detector array 400. In the illustrated embodiment, the detector array 400 is curved, and the detector elements of the detector array 400 are arranged to face an x-ray source located a fixed distance above along the y-axis from a center point 404 of the detector array 400, as indicated by dashed line 420 and reference axis 450.

[0036] The detector array 400 includes rails 440 between which collimating blades or plates 406 are positioned. The plates 406 collimate the x-rays 425 so that after collimating the x-rays 425, the beam impinges on a plurality of detector modules 402 of the detector array 400. The plurality of detector modules 402 may be positioned between the plates 406. Each detector module 402 may include a plurality of chiclets, each of which may include a plurality of detector elements or pixels. As an example, the detector array 400 may include 25 detector modules 402, each of which may include 16 chiclets, each of which may include 12 detector elements. Thus, each detector module 402 may include 192 detector elements, for a total of 4,800 detector elements in the detector array 400. In one example, the chiclet width is 8.74 mm, and each detector element has 100 pixels, each 365 μm wide in the X direction, with a 765 μm pitch in the Z direction. The small pixel size of 365 μm requires precise mechanical alignment that can be difficult to achieve, even if the position of the detector elements within each chiclet is controlled by robotic placement.

[0037] Returning to FIG. 2, because the reconstructed image is generated from the electrical signals generated by each detector element 202, the accuracy and / or quality of the reconstructed image depends on the position of the detector elements 202 within the detector array. Each detector element 202 is located within its respective detector array. Not all detector elements 202 are correctly positioned. If the actual positions of the detector elements 202 differ from their designed target positions (e.g., if the detector elements are misaligned), the accuracy of the reconstructed image may be reduced.

[0038] Detector elements 202 can be very small and difficult to align. For example, to achieve threshold image quality, it may be desirable to have alignment variations of less than 10% (e.g., less than 36.5 microns). If some of the detector elements 202 are misaligned by more than 10% relative to either or both of the x-axis and z-axis, the quality of the reconstructed image may be degraded.

[0039] In some examples, some of the detector elements 202 may become misaligned due to the manufacturing process or when the CT imaging system 200 is transported from a factory to a hospital. Some of the detector elements 202 may also become misaligned when a detector module of the detector array 108 is damaged and / or when a detector module of the detector array 108 is replaced with a new detector module. Misalignment may occur between detector modules (e.g., module-to-module alignment), or between different parts of a detector module (e.g., chiclet-to-chiclet alignment), or between sensors of a chiclet.

[0040] To ensure that threshold image quality is achieved, one or more detector arrays 108 of the CT imaging system 200 can be periodically calibrated. For example, one or more detector arrays 108 can be calibrated when the CT imaging system 200 is first installed in a medical facility, after a damaged module is replaced, after a predetermined time has elapsed, or at other times. During calibration, the position of each detector element 202 can be measured relative to a reference position provided by a wire phantom. Slight variations in the position of the detector element relative to the reference position can be detected, collected, and used to generate a calibration vector for each detector array 108. The calibration vector may include the measured position of the detector element or may include correction values ​​for the measured position of each detector element 202 of the detector array. The calibration vector can be used to perform mechanical alignment of the various detector elements of the detector array. Additionally or alternatively, during image reconstruction, the calibration vector can be applied to projection data acquired from the detector elements 202 to correct for variations in the position of the detector elements 202. Correcting the variations can improve image quality.

[0041] 3A shows an exemplary wire phantom assembly 300 used to calibrate one or more detector arrays 108. Wire phantom assembly 300 includes a wire phantom 302 and a table mount 320 that can be used to mount wire phantom 302 to a table (e.g., table 114) of a CT imaging system (such as CT imaging system 200 of FIG. 2).

[0042] The table mount 320 includes a mounting surface 334 that can be coupled to the surface of a table and a phantom holder 323 that can couple the wire phantom 302 to a desired position. The phantom holder 323 can adjustably position the wire phantom 302 a distance 325 below the underside of the table. For example, the distance 325 can be adjusted by bolting the phantom holder 323 to the table mount 320 at different locations 327 (e.g., holes in the phantom holder 323). The table mount 320 can include one or more rubber shims 322 coupled to the phantom holder 323, which can be positioned to make surface contact with the edge of the table. The rubber shims 322 can provide a firm and rigid coupling between the table mount 320 and the table. The base plate 336 of the wire phantom 302 can be coupled to the phantom holder 323 by one or more bolts (such as the first bolt 324 and the second bolt 330).

[0043] Briefly, referring to Figure 3B, table mount diagram 340 illustrates the alignment of wire phantom 302 and table mount 320 relative to table 342 of a CT imaging system. Mounting surface 334 may be aligned with underside 348 of table 342. For example, mounting surface 334 may be clamped to underside 348. The alignment of phantom holder 323 relative to underside 348 of table 342 with respect to the x-axis of the CT imaging system (e.g., according to reference axis 341) may be adjusted via first tilt screw 347. The alignment of phantom holder 323 relative to underside 348 of table 342 with respect to the y-axis of the CT imaging system may be adjusted via second tilt screw 349.

[0044] Returning to FIG. 3A , the wire phantom 302 includes a wire 304 disposed along the central axis of a tubular body 306, which is aligned with the z-axis of a CT imaging system. The tubular body 306 protects the wire 304 and may have a structure that maintains the straightness of the wire 304 and protects the wire 302. The wire phantom 302 includes a first end cap 308 disposed on a first side 314 (e.g., the gantry side) of the wire phantom 302 and a second end cap 312 disposed on a second side 316 (e.g., the table side) of the wire phantom 302. The wire 304 extends from a first inner surface of the first end cap 308 to a second inner surface of the second end cap 312. The first end cap 308, the second end cap 312, and the tubular body 306 have a diameter 307. For example, the diameter 307 may be 20 cm or another value.

[0045] The wire 304 can have a length 305 that is long enough to cover the entire detector array. For example, the length 305 can be 150 mm or longer. The wire 304 can include or be made of a highly attenuating material (such as steel or tungsten) that attenuates X-rays and provides a distinct signal when readout from the CT imaging system during a scan. During calibration, the position of each detector element can be measured relative to the position of the wire 304. A scan can be performed on the wire phantom 302. As the detector array is rotated around the table by a gantry (e.g., gantry 102), each detector element can detect the attenuation of X-rays by the wire 304 for different view angles.

[0046] The thickness or diameter of the wire 304 can be determined by the transaxial dimension of the detector element (e.g., pixel). To measure each pixel location reproducibly and reliably, a suitable diameter can be two to four times larger than the transaxial pixel dimension. In various embodiments, a suitable diameter is at least three times the transaxial dimension. In one embodiment, the transaxial dimension is 356 microns, and the diameter of the wire 304 is 1 mm (approximately three pixel widths). Thus, if the diameter of the wire 304 is approximately equal to three pixel widths, the attenuation of x-rays by the wire 304 during calibration can be detected by more than three pixels at a given field of view.

[0047] Wire 304 can be multi-threaded and / or braided, which makes it easier to achieve a threshold straightness than solid wire. For example, the threshold straightness can be a value that allows for a 30 micron variation from a straight line. In various embodiments, the minimum straightness standard can be maintained by tensioning wire 304 with springs 309. Additionally, wire 304 can be molded from a plastic resin, which further ensures that the minimum straightness standard is met.

[0048] The accuracy of the detector element positions measured during calibration depends on the accuracy of the angular alignment of the wires 304 relative to the table and gantry. In one example, accurate angular alignment is within a threshold of 3 milliradians (mrad) from being parallel to the central axis of the gantry. The accurate angular alignment can be controlled and / or adjusted via one or more tilt screws located on the phantom holder 323. Specifically, the position of the wires 304 relative to the table mount 320 can be adjusted via the first wire phantom tilt screw 326 and / or the second wire phantom tilt screw 332. The first wire phantom tilt screw 326 can adjust the vertical angle of the wire phantom 302 (e.g., in the y dimension indicated by the reference axis 301), and the second wire phantom tilt screw 332 can adjust the horizontal angle of the wire phantom 302 (e.g., in the x dimension). The wire 304 can be aligned or realigned parallel to the central axis of the gantry before or after each calibration and / or data analysis. Any remaining inaccuracies due to alignment accuracy of 3 milliradians or less can be corrected in a final calibration step by adjusting and fitting the wire itself.

[0049] 3C is a detector array diagram 360 illustrating an exemplary position of wire 304 relative to the detector array during calibration using wire phantom 302. Detector array diagram 360 includes an x-ray source 362 and a detector array 364. X-ray source 362 and detector array 364 may be mounted at fixed locations on a gantry (e.g., gantry 102). Center points 363 of x-ray source 362 and detector array 364 are aligned with the y-axis of the CT imaging system, and detector array 364 and x-ray source 362 may rotate about gantry isocenter 380. The gantry may rotate in the direction indicated by arrow 382.

[0050] Detector array 364 may have an arc shape with a first dimension in the x-axis direction of the CT imaging system and a flat shape with a second dimension in the z-axis direction of the CT system. During the calibration scan, the table of the CT system (not shown in FIG. 3C ) may be positioned in the z-axis direction, and wires 304 may be coupled to the end of the table to extend between detector array 364 and x-ray source 362. Thus, the position of wires 304 relative to detector array 364 and x-ray source 362 changes as the gantry rotates.

[0051] In the position shown in detector array diagram 360, x-rays 372 traveling toward detector element 368 are attenuated by wire 304 when they strike wire 304 at point 370 thereof, and thus detector element 368 detects a weaker x-ray 372. However, other x-rays emitted by x-ray source 362 from the same x-location (e.g., x-rays along the x-dimension indicated by arrow 384) can be detected by other detector elements positioned at the same x-location in detector array 364. Thus, x-rays 372 attenuated by wire 304 can be used to reveal a linear relationship between detector element 368, point 370 of wire 304, and x-ray source 362. Once this linear relationship is known, projection geometry can be used to determine the position of detector element 368 in detector array 364 and the angle of detector element 368 in the x- and z-axes relative to wire 304. How the positions of the detector elements 368 are calculated based on the attenuation of the x-rays 372 is described in more detail below with reference to FIGS.

[0052] 7, a wire position graph 700 shows an exemplary line 702 indicating the position of a wire of a wire phantom (such as wire 304 of wire phantom 302 of FIG. 3A) detected by multiple detector elements of a detector array over multiple view angles during one gantry rotation during a scan performed by a CT imaging system. The x-axis of wire position graph 700 shows frames corresponding to different view angles, and the y-axis of wire position graph 700 shows the position of a pixel (e.g., detector element) in the x-dimension.

[0053] Line 702 has the shape of a sinogram because the gantry rotates around the wire, as described above. During calibration, the wire phantom can be coupled to the table at a fixed position relative to the gantry. The fixed position can be outside the field of view (FOV) of the detector array to ensure that all detector elements of the detector array are calibrated. For example, the fixed position can be 50 cm outside the FOV. In other words, if the fixed position were assumed to be inside the FOV, some detector elements of the detector array that lie between the fixed position and the edge of the FOV would not detect the wire and would not be calibrated. Because the fixed position is outside the FOV, an upper portion 750 of line 702 and a lower portion 751 of line 702 are excluded from wire position graph 700.

[0054] For example, the first point 704 of the line 702 indicates that at the view angle corresponding to the 200th frame, a first attenuation of the x-rays due to the wire (referred to herein as wire attenuation) is most significantly detected by a first detector element at an x-position (350) of the detector array that includes the detector element. The first wire attenuation indicates the detector element at the x-position (e.g., 350) in the detector array.

[0055] As a second example, a second point 705 on line 702 shows that at a view angle corresponding to the 1350th frame, the second wire attenuation is most strongly detected by a first detector element at an x-position (350) of the detector array. The second wire attenuation is seen at the x-position (e.g., 350) of the detector element in the detector array. Each detector element in the detector array can detect the wire in two different frames during a full rotation.

[0056] Because the wire diameter is larger than the transverse diameter of a detector element in the detector array, the attenuation of x-rays by the wire can be detected by multiple adjacent detector elements. For example, if the wire diameter is 1 pixel and the detector element transverse diameter is 356 microns, three to five adjacent detector elements can record the attenuation. The attenuation is strongest at the center pixel of the three to five adjacent detector elements, and line 702 corresponds to the center pixel recording the wire attenuation as the gantry rotates. Wire attenuation is weaker in adjacent detector elements, which may produce wire shading 709 around line 702.

[0057] In wire position graph 700, wire shading 709 around line 702 has a horizontal extent 710, which is based on the number of view angles over which attenuation is recorded by a single detector element, and a vertical extent 720, which is based on the number of pixels that record attenuation at a given view angle.

[0058] For example, the first dashed line 706 indicates that in the 200th frame, the x-rays are most attenuated by the wire at the detector element at x-position 350 (see y-axis) corresponding to the first point 704. However, the x-rays are less attenuated at the detector elements adjacent in the x-direction, such as the detector elements corresponding to the first adjacent point 711 and the second adjacent point 713 (see y-axis).

[0059] Similarly, the second dashed line 712 indicates that a single detector element detects wire attenuation across multiple adjacent frames / view angles. As the gantry rotates, the detector element's position relative to the wire changes, and the attenuation detected by that detector element increases to a maximum attenuation and then decreases to zero. For example, the second dashed line 712 indicates that a detector element at x-position 1050 (see y-axis) records a first low attenuation in the 1650th frame, corresponding to point 718, a second maximum attenuation in the 1700th frame, corresponding to point 714 (e.g., a point located on line 702), and a third low attenuation in the 1750th frame, corresponding to point 719.

[0060] FIG. 8 shows a fitting diagram 800 illustrating how the exact x-position of a detector element in a detector array can be accurately determined from line 702 by fitting a function to measurements made by the detector element. For example, the function can be a Gaussian function, a parabolic function, or a different type of function. Fitting diagram 800 includes a wire position graph 802 similar to wire position graph 700, with an exemplary line 803 showing the position of a wire in a wire phantom in the form of a sinogram generated by wire attenuation measurements recorded at multiple detector elements / pixels. The y-axis of graph 802 shows the x-position of the pixel in the detector array, and multiple frames as the gantry rotates are shown on the x-axis, similar to FIG. 7. A close-up 804 shows a portion 812 of line 803, which shows the wire shading around line 803.

[0061] Graph 806 shows a plot 820 of measurements 822 collected at multiple pixels corresponding to portion 812, with plot 820 centered on a pixel at x-location (300) on the detector array indicated by dashed line 805. The intensity of the signal generated at the pixel is shown on the y-axis of graph 806, the x-location of the pixel is shown on the x-axis of graph 806, and each measurement shown on graph 806 represents the signal intensity recorded for the corresponding pixel, normalized between 0.0 and 1.0. The lowest signal intensity 824 is recorded for the pixel at x-location 300 where the x-rays are most attenuated by the wire (e.g., 0.3). The low signal intensity measurements at 828 and 829 (e.g., 0.4) indicate wire shading, where the x-rays are partially attenuated by the wire. Other adjacent pixels have less shading, but this shading does not extend beyond the two pixels on either side of the pixel at x-location 300.

[0062] To accurately determine the pixel location where the x-rays are most attenuated by the wire, a parabolic function 826 can be fitted to the measurements 822. Because the wire thickness (e.g., diameter) is larger (e.g., several times larger) than the transverse dimension of the detector element, multiple detector elements may record significant wire attenuation. The parabolic function 826 can be used to select a single detector element (e.g., pixel location) where wire attenuation is greatest. Using a wire that is larger than the transverse dimension of the detector element can improve pixel location accuracy relative to other wire phantoms that include thin wires. For example, alternative wire phantoms with thin wires may produce inconsistent and / or missing signals as the x-ray beam sweeps across the detector array. Alternatively, using a wire that is too thick may obscure too many detector elements, reducing pixel location accuracy. In various embodiments, the wire thickness can be selected based on the dimensions or characteristics of the CT system being calibrated. For example, in one embodiment, the selected thickness may be a function of a first distance between adjacent detector elements (e.g., detector pitch), a size of the focal spot of the CT system, and a second distance between the focal spot and the detector elements. In other embodiments, other or different variables may be considered.

[0063] In the attenuation signal weighted average, the statistical limitations of the X-ray count values ​​cause fluctuations in the measurements, which may not provide the required accuracy for the centroid of the wire position. The exemplary graph 806 shows fitting a parabolic function to reduce statistical or measurement-to-measurement fluctuations. Although the exemplary graph 806 shows fitting in the pixel direction to calculate the pixel centroid, fitting in the frame direction can also be performed to obtain the frame, depending on the final fitting model.

[0064] Once the single detector element with the greatest wire attenuation is selected, the projection geometry can be used to measure the position of the single detector element relative to the wire (e.g., the known position of the wire). Measuring the position of a single detector element is described in more detail below with reference to Figures 5A, 5B, and 9.

[0065] FIG. 5A is a flowchart illustrating an exemplary method 500 for using calibration data to improve the quality of images reconstructed using a CT system (such as the CT systems 100 and 200 described above). As described above, during calibration of a CT system, a calibration vector can be generated that includes the measured positions (or correction values) of each detector element of each detector array (e.g., detector arrays 108 and 400) of the CT imaging system. During subsequent scans of a subject (e.g., a patient) using the CT system, the calibration vector can be applied during image reconstruction to correct for misalignment of the detector elements. Specifically, processing of projection data during image reconstruction relies on the exact position of each detector element. However, the design target positions of detector elements, provided by, for example, detector design drawings based on the physical dimensions and characteristics of the CT imaging system, may differ from the measured actual positions of the detector elements. The measured positions, rather than the design target positions, can be used to reconstruct an image. Correcting the misalignment can improve the accuracy and quality of the reconstructed image. The method 500 may be stored as instructions in a non-transitory memory and executed by one or more processors of a computing device of a CT imaging system (such as computing device 216 of imaging system 200 of FIG. 2).

[0066] Method 500 begins at step 502, where method 500 includes performing a calibration of one or more detector arrays of a CT imaging system using a wire phantom. Performing a calibration of one or more detector arrays using a wire phantom is described in more detail with reference to FIG. 5B.

[0067] In step 504, method 500 includes storing the calibration vector generated during calibration of the CT imaging system in a memory of the CT imaging system (e.g., in the memory of computing device 216 of FIG. 2 ). The calibration vector may be retrieved and applied during processing of projection data acquired by the CT imaging system until a new calibration vector is generated during a subsequent calibration.

[0068] In step 506, the method 500 includes performing a scan on the subject. During the scan, projection data may be acquired by a CT imaging system. However, the projection data may contain pixel-level inaccuracies due to misalignment of detector elements (e.g., detector elements 202) arranged in a detector array of the CT imaging system.

[0069] In step 508, the method 500 includes applying a calibration vector during processing of the acquired projection data during image reconstruction. The calibration vector may include a measured position of each detector element in the detector array. During image reconstruction, signals generated by each detector element are processed based on the design target position of the detector element in the detector array. If the actual (e.g., measured) position of the detector element in the detector array differs from the design target position, the quality of the resulting image may be degraded. Therefore, the measured positions of the detector elements stored in the calibration vector may be used instead of the design target positions.

[0070] Alternatively, in some embodiments, the calibration vectors may include alignment correction values ​​to correct for detector element misalignment, and each position of each detector element in the projection data may be adjusted by applying the corresponding position correction value of the calibration vector. It should be understood that the calibration vectors referred to herein should not be confused with other different calibration vectors that are additionally applied, for example, to correct for variations in the signal from (properly aligned) detector elements.

[0071] In step 510, the method 500 includes displaying the reconstructed image on a display device (e.g., display device 232 of FIG. 2) of the CT imaging system and / or storing the reconstructed image in a memory of the CT imaging system, and the method 500 ends.

[0072] 5B illustrates a method 550 showing an exemplary procedure for calibrating a detector array of a CT imaging system using a wire phantom. Various steps of method 550 may be stored as instructions in a non-transitory memory and executed by one or more processors of a computing device of the CT imaging system (such as computing device 216 of imaging system 200 of FIG. 2). Some steps of method 550 are not executed by a computing device and may be executed manually, as described below.

[0073] Method 550 begins at step 552, where method 550 includes installing a wire phantom in a fixed position on a table (e.g., table 114) of a CT imaging system. The wire phantom can be installed manually, for example, by an operator of the CT imaging system. The wire phantom can be installed so that it is attached to the end or edge of the table, as described above in connection with FIG. 3B, and the wire (e.g., wire 304) of the wire phantom can extend into the space between the detector array and the x-ray source, as described above in connection with FIG. 3C. The fixed position of the wire can be outside the field of view of the detector array. As an example, the wire phantom can be attached to the table using the horizontal and vertical lasers of the gantry so that the wire is positioned 180 degrees below the isocenter of the gantry and 27 cm below the isocenter. A spirit level can be used on the wire tube to make it as parallel as possible to the Z axis of the CT imaging system. The center of the wire can be designated as the reference position of 0 mm on the Z axis.

[0074] In some examples, method 550 can include providing instructions to an operator to move the table to a specified height and / or to move the table to a specified position along the z-axis of the CT imaging system. For example, the operator can attach a wire phantom to the end or edge of the table and then issue the instructions via a display device (e.g., display device 232) of the CT imaging system.

[0075] In step 554, method 550 includes determining whether the wire is confirmed to be straight, e.g., whether it is within a threshold tolerance for straightness. In various embodiments, the wire can be confirmed to be straight based on input received from an operator. If there is variation in straightness, the wire cannot be used to calibrate the detector array. In one example, the threshold tolerance is 30 microns, and if any portion of the wire deviates from a straight line by more than 30 microns, the wire is not considered straight and cannot be used to calibrate the detector array. If the wire does not deviate from a straight line by more than 30 microns, the wire is confirmed to be straight and can be used to calibrate the detector array. If the wire phantom is new or has not been inspected recently, the wire straightness can be confirmed before proceeding with the calibration. For example, the wire straightness can be confirmed periodically (e.g., every six months) or after a predetermined number of scans (e.g., 600) have been performed.

[0076] If, in step 554, it is determined that the wire is not confirmed to be straight, method 550 proceeds to step 556. In step 556, method 550 includes performing an analysis of the wire's straightness. During the straightness analysis, rotational scans of the wire phantom can be performed, shifting the wire phantom to different positions along the z-axis. In other words, between each rotational scan, the position of the wire phantom in the z-axis direction can be adjusted in small increments (e.g., 5 mm). During each rotational scan, data can be collected from a small number of rows (e.g., 2-4 rows) in the center of the detector array, and data can be collected from the other rows of the detector array. The data collected from the small number of central rows can be used to determine whether the wire is within a threshold tolerance for straightness.

[0077] For example, with a wire phantom attached to the table, the position of the table in the z-axis direction can be adjusted so that the rotational plane (e.g., imaging plane) of the gantry in the x-y dimension intersects the wire of the wire phantom at a first point on the wire. The rotational plane can include a line (e.g., dashed line 420 in FIG. 4) along the y-axis of the CT imaging system between the X-ray source and the center point of the detector array (e.g., between X-ray source 362 and center point 363 in FIG. 3C), which is perpendicular to the z-axis. A first rotational scan can be performed to collect a first set of wire attenuation data from two to four central rows of the detector array. The table position can then be incrementally adjusted to a second specific point, and a second rotational scan can be performed to collect a second set of wire attenuation data from two to four central rows of the detector array. The table position is then adjusted incrementally to a third specific point and a third rotational scan is performed to collect a third set of wire attenuation data, and so on until data has been collected at multiple desired wire positions along the z-axis.

[0078] For example, the table is adjusted so that the imaging plane of the gantry intersects the wire at -55 mm relative to the zero (0) position at the center of the wire, and the table is adjusted along the z-axis by a predetermined length of 5 mm until it reaches a position of 55 mm relative to the zero (0) position, for a total of 23 scans, which is sufficient to cover the imaging FOV in the z-axis.

[0079] 6, a wire phantom alignment diagram 600 shows the position of a wire phantom 605 at three different times during wire attenuation data collection in one embodiment. The wire phantom 605 is coupled to a table mount 606 (e.g., table mount 320), which is coupled to a table 608 (e.g., table 114).

[0080] At a first time point 601, the wire phantom 605 is positioned so that the x-y rotation plane of the gantry, indicated by dotted line 607, intersects the wires 604 of the wire phantom 605 at a first point 614 of the wires 604. The imaging plane of the gantry may be perpendicular to the z-axis, as indicated by arrow 650 and reference axis 690, and includes an x-ray source 620 on a first side of the gantry, a center point 612 of a detector array 610 disposed on a second, opposite side of the gantry, and an isocenter point 611 about which the gantry rotates. The wire phantom 605 is offset from the isocenter point 611 so that a distance 615 between the wires 604 and the detector array 610 changes as the x-ray source 620 and detector array 610 rotate about the isocenter point 611. The first point 614 may be the center point of the wire 604, and the first point 614 may be a 0 mm landmark of the wire 604. At the first time point 601, the wire 604 is positioned within the FOV of the detector array 610.

[0081] At a second point in time 602, the position of the wire phantom 605 in the z-axis direction has shifted in a first positive direction indicated by a first arrow 622. The wire phantom 605 is positioned such that the imaging plane indicated by the dotted line 607 intersects the wire 604 of the wire phantom 605 at a second point 616 of the wire 604. The second point 616 may be a reference point of the wire 604 55 mm (e.g., 55 mm in the positive direction from the first point 614 (0 mm)). At the second point in time 602, the wire 604 is partially positioned outside the FOV of the detector array 610.

[0082] At a third point in time 603, the position of the wire phantom 605 in the z-axis direction has shifted in a second negative direction, indicated by a second arrow 624. The wire phantom 605 is positioned such that the imaging plane, indicated by the dotted line 607, intersects the wire 604 of the wire phantom 605 at a third point 618 of the wire 604. The third point 618 may be a reference point of the wire 604 −55 mm (e.g., 55 mm in the negative direction from the first point 614). At the third point in time 603, the wire 604 is positioned partially outside the FOV of the detector array 610.

[0083] To determine the straightness of the wire 604, data can be collected at multiple predetermined increments (e.g., a predetermined length of 5 mm) between the first point 614 and the third point 618, as described above. By collecting and comparing data at multiple predetermined increments, the straightness of the wire 604 can be accurately determined. For example, if a first set of wire attenuation data collected at a second time point 602 indicates a first position of a first portion of the wire 604, and a second set of wire attenuation data collected at a third time point 603 indicates a second position of a second portion of the wire 604, the deviation between the first and second positions indicates that the first portion is not aligned with the second portion, thereby inferring that the wire 604 is not straight. If the deviation is less than a threshold deviation (e.g., 30 microns), the wire 604 can be considered straight.

[0084] Returning to method 550, in step 558, method 550 includes determining whether the wire is straight according to the procedure described above with reference to Figure 6. If in step 558 it is determined that the wire is not straight, method 550 proceeds to step 560. In step 560, method 550 includes displaying instructions on a display device to the operator to fix or replace the wire, and method 550 returns to step 556. If in step 558 it is determined that the wire is straight, method 550 proceeds to step 562.

[0085] In step 562, method 550 includes performing an angular alignment analysis of the wire to determine the position of the wire relative to the z-axis of the CT imaging system (the table is aligned with the z-axis). In one embodiment, the wire is considered aligned with the z-axis if its deviation from an axis parallel to the z-axis is less than a threshold deviation. For example, the threshold deviation may be 3 mrad. In other words, if the wire deviates from an axis parallel to the z-axis by more than 3 mrad, the wire can be adjusted using the tilt adjustment screws (e.g., tilt screws 326 and 332 in FIG. 3A ), and alignment data collection and analysis can be repeated, as described in the next paragraph, until the deviation is less than 3 mrad.

[0086] A second set of rotational scans can be performed to perform a wire alignment analysis. The second set of rotational scans can be performed in a manner similar to that described above with reference to FIG. 6. However, because the wire is known to be straight, fewer scans can be performed. In one embodiment, two wire alignment scans can be performed. A first wire alignment scan of the wire phantom can be performed with the wire phantom positioned such that the imaging plane intersects the wire at a first distance in a first direction (e.g., arrow 622 in FIG. 6) from the center of the wire, i.e., the reference point (e.g., 0 mm). A second wire alignment scan of the wire phantom can be performed with the wire phantom positioned such that the imaging plane intersects the wire at a second distance in a second, opposite direction (e.g., arrow 624) from the center of the wire, i.e., the reference point. The second distance can be equal to the first distance. The first and second distances can be separated by a threshold distance. For example, the threshold distance can be 80 mm.

[0087] For example, the first distance and the second distance can both be 40 mm. First, the table can be adjusted to position the wire so that the imaging plane intersects the wire at a position 40 mm away from the center of the wire (i.e., the reference point), and a first rotational scan can be performed. Then, the table can be adjusted to position the wire so that the imaging plane intersects the wire at a position −40 mm away from the center of the wire (i.e., the reference point), and a second rotational scan can be performed. In each of the first and second rotational scans, wire attenuation data can be acquired from the central two to four rows of detector elements of the detector array. The angular alignment analysis can be performed faster than the straightness analysis because it has already been confirmed that the wire is straight.

[0088] By collecting and comparing data from the first and second rotational scans, the angle of the wire 604 relative to the z-axis can be accurately determined. For example, if a first set of wire attenuation data collected during a first rotational scan indicates a first position of the wire 604, and a second set of wire attenuation data collected during a second rotational scan indicates a second position of the wire 604, then a deviation between the first and second positions indicates that the wire is misaligned. If the deviation is less than a threshold deviation amount (e.g., 3 mrad), the wire can be considered aligned with the z-axis.

[0089] In step 564, method 550 includes determining whether the wire is aligned with the z-axis, according to the procedure described above. If it is determined that the wire is not aligned with the z-axis (e.g., if the deviation is greater than a threshold deviation amount), method 550 proceeds to step 566. In step 566, method 550 includes displaying instructions on a display device to readjust the wire. In various embodiments, the wire readjustment can be performed by an operator adjusting one or more tilt screws of the wire phantom. For example, a vertical tilt screw (e.g., tilt screw 326 of phantom holder 323) located in the phantom holder of the wire phantom can adjust the wire with respect to the y-axis, and a horizontal tilt screw (e.g., tilt screw 332 of phantom holder 323) can adjust the wire with respect to the x-axis. After the wire is readjusted, method 550 can return to step 562 and perform the wire angular alignment analysis again. In some cases, the wire is readjusted and the wire angular alignment analysis is performed multiple times to ensure the wire is aligned with the z-axis.

[0090] If, in step 564, it is determined that the wire is aligned with the z-axis, method 550 proceeds to step 568. In step 568, method 550 includes performing a calibration scan to measure the precise positions of the individual detector elements (detector elements 202) of the detector array relative to the wire. Because the position of the wire is known and fixed, the precise positions of the individual detector elements can be measured by performing a full rotation, fitting a function to the sinogram generated by the projection data, and using a projection geometry based on the physics of the CT system, as described above with reference to Figures 7 and 8.

[0091] Referring to Figure 9, a projection geometry diagram 900 illustrates several variables used to calculate the position of the wire. The projection geometry diagram 900 illustrates an exemplary CT system 902 (e.g., CT system 100 or imaging system 200) including an x-ray source 904 and a detector array 906, which is a non-limiting example of the detector array 400 of Figure 4. The x-ray source 904 and the detector array 906 may be coupled to a gantry (e.g., gantry 104) that rotates in a direction 901 about an isocenter (point) 914, which is midway between the x-ray source 904 and a center point 907 of the detector array 906 on an isocenter line 905. The detector array 906 includes multiple detector modules 908, each of which includes multiple pixels 912. The pixels 912 may be arranged in slices or chiclets 910. Each pixel 912 of the detector array 906 can be identified in the x-dimension by a distance 909 along the detector array 906 from the center point 907 based on a reference axis 990, and in the z-dimension by the row of the pixel 912. The wires of the wire phantom are positioned at wire positions 916 within the CT system 902.

[0092] 9 shows the CT system at a point where the gantry has rotated from an initial position (shown by line 926) by a gantry rotation angle 922 (β) relative to line 926. The wire position 916 (relative to the x-ray source 904 and / or detector array 906) can be measured based on the distance R (shown by line 917) between the wire position 916 and the isocenter 914 and the angle 918 (φ) between line 917 and line 926.

[0093] At the time shown in FIG. 9 , an X-ray beam emitted by X-ray source 904 is detected by detector array 906. However, the X-ray beam on its trajectory, indicated by line 919, strikes the wire at wire position 916, and pixel 920 fails to detect the X-ray beam. Line 919 is at angle 924 (γ) relative to isocenter line 905. Therefore, the angular position γ of pixel 920 relative to isocenter line 905 can be expressed as a function of gantry rotation angle β based on R and φ, where R and φ are the wire positions measured in the previously described wire alignment step. R is the wire distance from the isocenter, and φ is the wire angle from isocenter line 905. SID is the distance from source 904 to isocenter 914, which is known by engineering design. β 922 is the gantry rotation angle relative to the gantry's vertical Y-axis 926.

number

number

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[0094] In step 570, method 500 may optionally include generating alignment correction values ​​from the measured positions of the detector elements. The alignment correction values ​​may be values ​​based on differences between the measured positions of the detector elements and their design target positions.

[0095] In step 572, method 500 may include storing the measured positions (or alignment corrections) of the detector elements in a calibration vector for use in subsequent scans of the subject. Method 550 then ends. In other words, a calibration vector may be generated that includes the measured positions of each detector element or corrections for adjusting the position of the detector element to correct for the detected and quantified misalignment. The calibration vector may be used in method 500 of FIG. 5A, as described above, to improve the quality of reconstructed images of the subject during subsequent scans.

[0096] Thus, a method for correcting misalignment of detector elements in a detector array of a CT system is provided. This method can be used when mechanical calibration techniques are not feasible due to the size of the detector elements. This method uses a wire phantom, which, unlike other wire phantoms, contains wires thicker than the detector elements. In an initial step, the wire phantom is attached to the table of the CT system, and the straightness and tilt angle of the wires are confirmed according to the procedures described herein. Once the straightness and proper alignment of the wires are confirmed, the position of each detector element in the detector array relative to the wires is measured. The measured positions are stored in a calibration vector, which can be applied during image reconstruction during subsequent scans of the subject to correct for misalignment.

[0097] By using measured detector element positions rather than the target positions of the design of an engineering model of the detector array, detector element misalignment can be corrected during scanning of the subject, improving the quality of the reconstructed image obtained from the scan. An advantage of using measured detector element positions is that detector element misalignment can be adjusted without relying on mechanical alignment techniques. As a result, the method can be applied to detector arrays containing smaller detector elements than can be aligned using mechanical alignment techniques. Furthermore, CT system calibration time, resource allocation, and cost are reduced, making the CT system more efficient. This improves CT system functionality, reduces downtime during calibration, and enables faster and more accurate processing. Furthermore, by applying the calibration method described herein, CT systems can produce higher-quality images, resulting in more accurate diagnoses and more desirable patient outcomes.

[0098] The technical effect of using a wire phantom to calibrate the positions of detector elements in a detector array of a CT system is to improve the quality of images reconstructed using the CT system.

[0099] The present disclosure also supports a method for a computed tomography (CT) system. The method includes measuring positions of detector elements of a detector array of the CT system during calibration of the CT system using wires of a wire phantom coupled to a table of the CT system during a rotational scan performed using the CT system, applying the measured positions of the detector elements instead of design target positions of the detector elements during image reconstruction from projection data acquired by the CT system during a subsequent scan performed on a subject using the CT system, and displaying the reconstructed image on a display device of the CT system. In a first embodiment of the method, the design target positions of the detector elements are provided by a detector design drawing of the CT system and / or a dimensionally based model of the CT system. A second embodiment of the method optionally includes the first embodiment and mechanically adjusts positions of detector elements of the detector array using the measured positions of the detector elements.

[0013] A third embodiment of the method optionally includes one or both of the first and second embodiments, and further includes, prior to measuring positions of the detector elements relative to a position of the wire, performing at least one of a first analysis of the straightness of the wire and a second analysis of an inclination angle of the wire relative to a z-axis of the CT system parallel to a longitudinal direction of the table. A fourth embodiment of the method optionally includes one or more or each of the first to third embodiments, and further includes performing a first analysis of the straightness of the wire, incrementally adjusting a position of the wire along the z-axis during each rotational scan of the multiple rotational scans, measuring positions of detector elements located in a small number of rows in a central portion of the detector array relative to the wire during each rotational scan, and not measuring positions of detector elements located outside the small number of rows in the central portion of the detector array relative to the wire, and comparing the measured positions of the detector elements in each rotational scan of the multiple rotational scans to determine the straightness of the wire.A fifth example of the method optionally includes one or more of the first to fourth examples, and further includes performing the second analysis of the tilt angle of the wire relative to a z-axis of the CT system by performing a first rotational scan with the wire positioned at a first position on the z-axis, performing a second rotational scan with the wire positioned at a second position on the z-axis, the second position and the first position being separated by a threshold distance, measuring positions of detector elements positioned in a small number of rows in a central portion of the detector array relative to the wire during each rotational scan, and not measuring positions of detector elements positioned outside the small number of rows in the central portion of the detector array relative to the wire, and comparing the measured positions of the detector elements in the first rotational scan with the measured positions of the detector elements in the second rotational scan to determine the tilt angle of the wire. A sixth example of the method optionally includes one or more of the first to fifth examples, and the small number is a number between 2 and 4. A seventh embodiment of the method optionally includes the or each of the first to sixth embodiments, and further includes, in response to a result of the second analysis indicating that the tilt angle of the wire is outside a second threshold deviation, displaying instructions to an operator of the CT system to adjust the tilt angle of the wire by one or more tilt screws disposed on a phantom holder used to couple the wire phantom to the table. An eighth embodiment of the method optionally includes the or each of the first to seventh embodiments, wherein the wire phantom is coupled to an edge of the table such that the position of the wire is outside a field of view of the detector array. A ninth embodiment of the method optionally includes the or each of the first to eighth embodiments, wherein a diameter of the wire is at least three times the transverse dimension of the detector element.In a tenth embodiment of the method, optionally including one or more of the first to ninth embodiments, measuring the positions of the detector elements relative to the position of the wire further includes measuring the positions of the detector elements based on known physical characteristics and projection geometry of the CT system. In an eleventh embodiment of the method, optionally including one or more of the first to tenth embodiments, measuring the positions of the detector elements relative to the position of the wire further includes fitting one of a parabolic function and a Gaussian function to measurements of signals generated at adjacent detector elements including the detector element. In a twelfth embodiment of the method, optionally including one or more of the first to eleventh embodiments, further includes calculating a respective alignment correction value for each detector element of the detector array based on the measured positions of the detector elements, storing the calculated alignment correction values ​​for all of the detector elements in a calibration vector, and correcting the positions of all of the detector elements of the detector array using the calibration vector during image reconstruction.

[0100] The present disclosure also supports a computed tomography (CT) system including a processor and non-transitory memory containing instructions that, when executed, cause the processor to perform a first series of rotational scans of a wire phantom to determine straightness of wires in the wire phantom, perform a second series of rotational scans of the wire phantom to determine a tilt angle of the wires, and, in response to the straightness of the wires being less than or equal to a first threshold deviation and the tilt angle of the wires being less than or equal to a second threshold deviation, measure a position of each detector element of a detector array of the CT system relative to a position of the wires during a calibration scan performed using the CT system, apply the measured position of each detector element during reconstruction of images from projection data acquired by the CT system during subsequent scans performed using the CT system, and display the reconstructed images on a display device of the CT system. In a first embodiment of the system, performing the first series of rotational scans of the wire phantom to determine the straightness of the wire and performing the second series of rotational scans of the wire phantom to determine the tilt angle of the wire comprises adjusting the position of the table of the CT system on which the wire phantom is mounted along a z-axis of the CT system parallel to a longitudinal direction of the table during each rotational scan of the first series of rotational scans and the second series of rotational scans, so that an imaging plane of a gantry of the CT system intersects the wire at a different position along the length of the wire during each rotational scan. A second embodiment of the system optionally includes the first embodiment, wherein the wire phantom is coupled to the table such that the wire is located outside the FOV of the detector array.In a third embodiment of the system, optionally including one or both of the first and second embodiments, measuring the position of each detector element of a detector array of the CT system relative to the position of the wire further includes collecting energy measurements at each detector element of the detector array during a full rotation of the calibration scan, fitting a function to the energy measurements collected during each frame of the full rotation at a view angle corresponding to the frame, selecting detector elements of the detector array based on the function, and measuring the position of the selected detector elements relative to the position of the wire. In a fourth embodiment of the system, optionally including one or more or each of the first to third embodiments, measuring the position of the selected detector elements relative to the position of the wire further includes measuring the position based on known physical characteristics and projection geometry of the CT system.

[0101] The present disclosure also supports a wire phantom assembly for calibrating a computed tomography (CT) system. The wire phantom assembly includes a wire phantom and a phantom holder, the phantom holder being attachable to a table of the CT system and including one or more tilt screws for adjusting the alignment of the wire phantom relative to the phantom holder. The wire phantom includes a tubular body surrounding a braided wire molded from a plastic resin, the braided wire including a material with high x-ray attenuation, and the braided wire having a diameter at least three times the transverse dimension of a detector element of the CT system. In a first embodiment of the system, the straightness of the wire is maintained by spring tension.

[0102] When introducing elements of various embodiments of the present disclosure, the articles “a,” “an,” and “the” are intended to mean that there are one or more of the element. The terms “first,” “second,” and the like do not denote order, quantity, or importance, but rather are used to distinguish one element from another. “Comprising,” “including,” “having,” and the like are intended to be inclusive and mean that there may be additional elements other than the listed elements. As used herein, terms such as “connected,” “coupled,” and the like refer to one object (e.g., a material, element, structure, member, etc.) being connected to or coupled to another object, regardless of whether the one object is directly connected to or coupled to the other object, or whether there are one or more intervening objects between the one object and the other object. In addition, it should be understood that references to “one embodiment” or “an embodiment” of the present disclosure are not intended to be interpreted as excluding the existence of additional embodiments that also incorporate the referenced features.

[0103] In addition to the modifications described above, many other variations and alternative structures may be devised by those skilled in the art without departing from the spirit and scope of the present description, and the claims are intended to cover such modifications and structures. Thus, while the above information has been described in particular detail with respect to what are presently considered to be the most practical and preferred embodiments, it will be apparent to those skilled in the art that many modifications, including but not limited to, in form, function, method of operation, and use, are possible without departing from the principles and concepts described herein. Furthermore, the examples and embodiments herein are meant to be merely illustrative in all respects and should not be construed as limiting in any manner. [Explanation of symbols]

[0104] 100 CT System 102 Gantry 104 X-ray source 106 X-ray radiation beam 108 detector array 110 Image Processing Unit 112 Subject 114 Tables 200 CT Imaging System 202 detector element 204 specimens 206 Center of rotation 208 Control Mechanism 210 X-ray controller 212 Gantry motor controller 214 Data Acquisition System (DAS) 216 Computing Devices 218 Mass storage 220 Operator Console 226 Table Motor Controller 230 Configurator 232 Display device 300 Wire Phantom Assembly 301 Reference axis 302 Wire Phantom 304 wire 305 length 306 Tubular body 307 diameter 308 First End Cap 309 Spring 312 Second End Cap 314 First Side 316 Second Side 320 Table mounting part 322 Rubber Shim 323 Phantom holding part 324 First Bolt 325 distance 326 Inclined Screw 327 position 330 Second Bolt 332 Inclined Screw 334 Mounting surface 336 base plate 341 Reference axis 342 Table 347 First Inclined Screw 348 Bottom surface 349 Second Inclined Screw 350x position 362 X-ray source 363 center point 364 detector array 368 detector elements 370 points 372 X-ray 380 Isocenter 382 Arrow 384 Arrow 400 detector array 402 Detector Module 404 Center point 406 Plate 420 dashed line 425 X-ray 440 Rail 450 Reference axis 500 ways 502 steps 504 steps 506 steps 508 steps 510 steps 550 method 552 steps 554 steps 556 steps 558 steps 560 steps 562 steps 564 steps 566 steps 568 steps 570 steps 572 steps 601 First Time 602 Second Time 603 Third Time 604 Wire 605 Wire Phantom 606 Table mounting part 607 dotted line 608 Table 610 detector array 611 Isocenter point 612 center point 614 First Point 615 distance 616 Second Point 618 Third Point 620 X-ray source 622 First Arrow 624 Second Arrow 650 Arrow 690 Reference axis 700 Wire Position Graph 702 Line 704 First Point 705 Second Point 706 First dashed line 709 Wire Shading 710 horizontal range 711 first adjacent point 712 Second dashed line 713 Second adjacent point 714 points 718 points 719 points 720 vertical range 750 Upper part 751 Lower part 802 graphs 803 Line 805 dashed line 806 graphs 812 part 812 part 820 plots 822 measurements 824 signal strength 826 Parabolic Function 901 direction 902 CT System 904 X-ray source 905 Isocenter Line 906 detector array 907 center point 908 Detector Module 909 distance 910 Chiclet 912 pixels 914 Isocenter 916 Wire Position 917 line 918 angle 919 line 920 pixels 922 rotation angle 924 angle 990 Reference axis

Claims

1. 1. A method for a computed tomography (CT) system, comprising: During calibration of the CT system, measuring (568) positions of detector elements of a detector array of the CT system using wires of a wire phantom coupled to a table of the CT system during a rotational scan performed using the CT system; During a subsequent scan performed on the subject using the CT system, applying the measured positions of the detector elements rather than the design target positions of the detector elements during image reconstruction from projection data acquired by the CT system (508); and Displaying the reconstructed image on a display device of the CT system (510). A method comprising:

2. The method of claim 1 , wherein the design target positions of the detector elements are provided by a detector design drawing of the CT system and / or a dimensionally based model of the CT system.

3. The method of claim 1 , wherein the measured positions of the detector elements are used to mechanically adjust positions of detector elements of the detector array.

4. 2. The method of claim 1, further comprising, before using the wire to measure the position of the detector element, performing at least one of a first analysis (556) of the straightness of the wire and a second analysis (562) of the tilt angle of the wire relative to a z-axis of the CT system that is parallel to the longitudinal direction of the table.

5. Performing a first analysis (556) of the wire straightness includes: performing a plurality of rotational scans and incrementally adjusting the position of the wire along the z-axis during each rotational scan of the plurality of rotational scans; measuring positions of detector elements disposed in a small number of rows in a central portion of the detector array relative to the wire during each rotational scan, and not measuring positions of detector elements disposed outside the small number of rows in the central portion of the detector array relative to the wire; and comparing the measured positions of the detector elements in each rotational scan of the plurality of rotational scans to determine the straightness of the wire; The method of claim 4 further comprising:

6. When the second analysis is performed, performing the second analysis (562) of the tilt angle of the wire relative to the z-axis of the CT system includes: performing a first rotational scan with the wire positioned at a first position on the z-axis; performing a second rotational scan with the wire positioned at a second position on the z-axis, the second position being separated from the first position by a threshold distance; measuring positions of detector elements disposed in a small number of rows in a central portion of the detector array relative to the wire during each rotational scan, and not measuring positions of detector elements disposed outside the small number of rows in the central portion of the detector array relative to the wire; and comparing the measured positions of the detector elements in the first rotational scan with the measured positions of the detector elements in the second rotational scan to determine a tilt angle of the wire; The method of claim 4 further comprising:

7. The method of claim 6, wherein the small number is between two and four.

8. 5. The method of claim 4, further comprising: in response to a result of the second analysis indicating that the tilt angle of the wire is outside a second threshold deviation, displaying (566) instructions to an operator of the CT system to adjust the tilt angle of the wire by one or more tilt screws disposed in a phantom holder used to couple the wire phantom to the table.

9. The method of claim 1 , wherein the wire phantom is coupled to the table at a fixed position, the fixed position of the wire being outside the FOV of the detector array.

10. The method of claim 1 , wherein the diameter of the wire is at least three times the transverse axial extent of the detector element.

11. The method of claim 1 , wherein measuring the positions of the detector elements using the wires further comprises measuring the positions of the detector elements based on known physical characteristics and projection geometry of the CT system.

12. 10. The method of claim 1, wherein measuring the position of the detector element using the wire further comprises fitting one of a parabolic function and a Gaussian function to measurements of signals produced at a plurality of adjacent detector elements, including the detector element.

13. calculating (570) a respective alignment correction value for each detector element of the detector array based on the measured positions of the detector elements; storing the calculated alignment correction values ​​for all of the detector elements in a calibration vector (572); and correcting the positions of all detector elements of the detector array using the calibration vector during image reconstruction (508). The method of claim 1 further comprising:

14. A computed tomography (CT) system (100, 200) including a processor and non-transitory memory, the non-transitory memory including instructions that, when executed, cause the processor to: performing a first series of rotational scans of a wire phantom (302) to determine the straightness of wires (304) of said wire phantom (302); performing a second series of rotational scans of the wire phantom (302) to determine the tilt angle of the wire (304); In response to the straightness of the wire (304) being equal to or less than a first threshold deviation amount and the tilt angle of the wire (304) being equal to or less than a second threshold deviation amount, While a calibration scan is being performed using the CT system (100, 200), measuring the position of each detector element (912) of a detector array (906) of the CT system (100, 200) relative to the position of the wire (304); and While a subsequent scan is being performed using the CT system (100, 200), applying the measured positions of each detector element (912) during reconstruction of an image from projection data acquired by the CT system (100, 200); and Displaying the reconstructed image on a display device of the CT system (100, 200). A CT system that performs the above.

15. 15. The CT system of claim 14, wherein performing the first series of rotational scans of the wire phantom to determine the straightness of the wire and performing the second series of rotational scans of the wire phantom to determine the tilt angle of the wire comprises adjusting the position of the table of the CT system on which the wire phantom is mounted along a z-axis of the CT system parallel to a longitudinal direction of the table during each rotational scan of the first series of rotational scans and the second series of rotational scans, so that an imaging plane of a gantry of the CT system intersects the wire at a different position along the length of the wire during each rotational scan.

Images