X-ray detector with rectangular pixels

Abstract

The invention relates to an X-ray detector with rectangular pixels. Systems and methods for a radiation detector having rectangular pixels are provided herein. In one example, an x-ray imaging system (100) includes a pixel array of a flat panel detector (18) including a plurality of pixels having a rectangular pixel pitch, the plurality of pixels being arranged in pairs (300), where each pixel of the plurality of pixels is configured to generate a respective image data signal, where in a low dose application, the image data signal is transmitted to the panel detector (18). The TFT control lines (320, 322) of the pixels in each pixel pair are simultaneously energized to generate a signal having an effective pixel pitch that is twice the rectangular pixel pitch, and in a high dose application, the TFT control lines (320, 322) of the pixels in each pixel pair are sequentially energized to generate a signal having an effective pixel pitch that is twice the rectangular pixel pitch. And the detector is translated by an effective pixel pitch that is half the rectangular pixel pitch during image acquisition.

Description

TECHNICAL FIELD

[0001] Embodiments of the subject matter disclosed herein relate to medical imaging, and more particularly, to an x-ray detector with rectangular pixels. BACKGROUND

[0002] Solid state radiation imagers are used to image non-optical radiation, such as x-rays, and high energy nuclear radiation, such as gamma rays. Typically, incident radiation is detected in a solid state imager by a process of absorption of the incident radiation in a scintillator, producing optical photons. Photodetectors, such as photodiodes, disposed in an array adjacent to the scintillator detect the optical photons. The location in the array of photodetectors that detects the light and the intensity of the signal generated by the photodetector are processed to display and analyze the incident radiation. Alternatively, the incident radiation can be directly absorbed into a photosensitive element that converts the energy of the incident radiation into mobile charge particles.

[0003] One typical use of solid state radiation imaging systems is medical imaging, where radiation that passes through or is emitted from a patient's body is used to visualize objects or materials within the body. Medical imaging devices preferably exhibit high spatial frequency response and high signal-to-noise ratio, while maintaining a feasible manufacturing and operating expenditure for various types of scans. SUMMARY

[0004] In one example, an x-ray imaging system includes: an array of photodetector pixels of a detector, the array of photodetector pixels including a plurality of pixels having a rectangular pixel pitch, wherein the plurality of pixels are arranged in pairs, and each pixel of the plurality of pixels is configured to generate a respective image data signal via a thin film transistor (TFT), a TFT control line, and a data readout line, wherein in a low dose application, the TFT control lines of the pixels of each pair of pixels are simultaneously energized to generate the respective image data signal having a first effective pixel pitch that is twice the rectangular pixel pitch, and in a high dose application, the TFT control lines of the pixels of each pair of pixels are sequentially energized to generate the respective image data signal having the rectangular pixel pitch; and a controller coupled to the array of photodetector pixels to receive the respective image data signals generated by the plurality of pixels, wherein, in the high dose application, the controller is configured to translate between a first position and a second position to generate two image data signals for each pixel of the plurality of pixels, and the controller is configured to combine the two image data signals to generate an image data signal having a second effective pixel pitch that is half the rectangular pixel pitch.

[0005] It is to be understood that the above brief description is provided to introduce in simplified form a selection of concepts that are further described in the detailed description. This brief description does not identify key or essential features of the claimed subject matter nor does it BRIEF DESCRIPTION OF DRAWINGS

[0006] The present application will be better understood by reading the following description of non-limiting embodiments, with reference to the appended drawings, in which:

[0007] Figure 1 A pictorial view of an imaging system is shown in accordance with one embodiment;

[0008] Figure 2 A pictorial view of an exemplary imaging system is shown in accordance with one embodiment;

[0009] Figure 3 An exemplary pair of rectangular pixels in a pixel array is shown in accordance with one embodiment;

[0010] Figure 4 A relative position of an x-ray detector compared to a pair of rectangular pixels is shown in accordance with one embodiment;

[0011] Figure 5 A flowchart of a method for generating a low dose x-ray image is shown in accordance with one embodiment; and

[0012] Figure 6 A flowchart of a method for generating a high dose x-ray image is shown in accordance with one embodiment. DETAILED DESCRIPTION

[0013] The following description relates to various embodiments of medical imaging. In particular, systems and methods for x-ray detectors having rectangular pixels are provided. In Figure 1 Examples of x-ray systems are shown in Figure 2 The exemplary x-ray system shown in Figure 1 and Figure 2 The exemplary x-ray system shown is a mammography system, but it is to be understood that other x-ray based systems, such as computed tomography (CT) systems, etc., can also be applied without departing from the scope of the present disclosure. Figure 3 A diagram of a pair of rectangular pixels in accordance with the present disclosure is shown in Figure 4 A schematic diagram of an x-ray detector positioned relative to a pair of rectangular pixels is shown in

[0014] Figure 5 and Figure 6Methods for operating an x-ray machine to acquire low dose x-ray images and high dose x-ray images are shown in the figures.

[0015] Current flat panel x-ray detectors, such as those used in mammography, CT, and other types of medical imaging, utilize x-ray detectors having an array of pixels. The pixel size is statically fixed, for example, 100 micrometers (pm) x 100 pm or 50 pm x 50 pm. However, the pixel size impacts the spatial resolution and the signal-to-noise ratio. For example, when the detector array includes larger pixels (e.g., 100 pm x 100 pm pixels), the high dose application images can have a lower spatial resolution, and when the detector array includes smaller pixels (e.g., 50 pm x 50 pm), the high dose application images can have a high spatial resolution. However, when the detector array includes smaller pixels, the low dose application images can have a lower signal-to-noise ratio, and when the detector array includes larger pixels, the low dose application images can have a higher signal-to-noise ratio. Thus, for low dose applications, a large pixel size (e.g., 100 pm x 100 pm) can be preferred, but in high dose applications with large size, the spatial resolution can be impacted. Thus, for high dose applications, a small pixel size (e.g., 50 pm x 50 pm) can be preferred, but in low dose applications with small pixel size, the signal-to-noise ratio can be impacted. To overcome this, high dose applications and low dose applications can be performed on separate imaging systems. However, this is inconvenient for hospital systems, both in terms of scheduling of imaging studies, and in terms of footprint.

[0016] Further, a widely used photodiode for x-ray imaging is amorphous silicon (a-Si). However, a-Si photodiodes can be used with detector arrays having a minimum pixel size of approximately 75 x 75 pm. Thus, the fabrication of a-Si photodiodes does not allow for smaller pixel sizes, such as 50 pm x 50 pm. However, other photodiodes, such as complementary metal-oxide-semiconductor (CMOS), which can be used with detector arrays having smaller pixel sizes, are less readily available and / or less cost effective.

[0017] Accordingly, provided herein are methods and systems for x-ray systems that address, at least in part, these issues, where the x-ray systems include pairs of rectangular x-ray detector pixels. For example, the pixels can have a first dimension (e.g., height) of 100 pm and a second dimension (e.g., width) of 50 pm. When used for low dose applications, pairs of adjacent rectangular pixels can be powered simultaneously and can be read out together. Accordingly, the pair of rectangular pixels can be read out as a pixel having an effective larger (e.g., 100 pm x 100 pm) pixel size to generate a single image. Accordingly, the rectangular pixels provide a high signal-to-noise ratio in low dose applications. When used for high dose applications, the pixels in a pair of rectangular pixels can be powered sequentially and read out individually to generate two images having rectangular pixels (e.g., 100 pm x 50 pm). Additionally, the detector can be translated along the first dimension by 50 pm such that the two generated images have a positional difference of 50 pm. The two images are then combined into one image having an effective small pixel size (e.g., 50 pm x 50 pm) using image processing techniques. In this way, with a single x-ray system, the pixel size can be optimized for both low dose applications and high dose applications. Moreover, the rectangular pixels can have dimensions that are compatible with a-Si photodiodes, providing for easier manufacturing.

[0018] Referring to Figure 1 FIG. 1 shows a digital mammography system 100 including an x-ray system 10 for performing a mammography procedure, in accordance with embodiments of the present disclosure. The x-ray system 10 can be used to acquire one or more types of images during one or more acquisitions. It should be understood that the digital mammography system 100 described is an example of an x-ray imaging system, and other types of x-ray imaging systems using flat panel detectors, such as CT systems, are also applicable. Figure 1 The digital mammography system 100 described is an example of an x-ray imaging system, and other types of x-ray imaging systems using flat panel detectors, such as CT systems, are also applicable.

[0019] The x-ray system 10 includes a support structure 42 to which the radiation source 16, the radiation detector 18, and the collimator 20 are attached. The radiation source 16 is housed within a gantry 15 that is movably coupled to the support structure 42. In particular, the gantry 15 can be mounted to the support structure 42 such that the gantry 15, including the radiation source 16, can be rotated relative to the radiation detector 18 about an axis 58. The range of rotation of the gantry 15 housing the radiation source 16 indicates a rotation in either direction about the axis 58 up to a desired angle. For example, the range of rotation of the radiation source 16 can be -0 to +0, where 0 can be such that the range of angles is a limited range of angles that is less than 360 degrees. An example x-ray system can have a range of angles of ±11 degrees, which can allow the gantry to be rotated from -11 degrees to +11 degrees about the axis of rotation of the gantry (i.e., rotation of the radiation source). The range of angles can vary depending on manufacturing specifications. The range of angles for a digital mammography system can be approximately ±11 degrees to ±60 degrees, depending on manufacturing specifications.

[0020] Radiation source 16 is directed toward a volume or object to be imaged and is configured to emit radiation rays at a desired time to acquire one or more images. Radiation detector 18 is configured to receive radiation rays via surface 24. Detector 18 can be any flat panel detector. Detector 18 can include an array of photosensor pixels optically coupled to a scintillator (e.g., of surface 24). The scintillator can be disposed to receive incident radiation rays, such as x-rays. Optical photons generated in the scintillator in response to the incident radiation are transferred to the array of photosensor pixels, where the light is detected and corresponding image data signals are generated by the photosensors in the array. Collimator 20 is disposed adjacent to radiation source 16 and is configured to adjust the subject’s irradiation zone. Further, detector 18 can be configured to linearly translate using a fast piezoelectric linear actuator.

[0021] In some embodiments, system 10 can also include a patient shield 36 mounted to radiation source 16 via a shield rail 38 so that a body part of a patient (e.g., a head) is not directly under the radiation. System 10 can also include a compression paddle 40 that is movable up and down relative to the support structure along a vertical axis 60. Thus, compression paddle 40 can be adjusted to be positioned closer to radiation detector 18 by moving the compression paddle 40 down toward the detector 18, and the distance between detector 18 and compression paddle 40 can be increased by moving the compression paddle up along vertical axis 60 away from the detector. Movement of compression paddle 40 can be adjusted by a user via a compression paddle actuator (not shown) included in x-ray system 10. Compression paddle 40 can hold a body part, such as a breast, in place against surface 24 of radiation detector 18. Compression paddle 40 can compress the body part and hold it in place while optionally providing an aperture to allow insertion of a biopsy needle, such as a core needle or a vacuum assisted core needle. In this way, compression paddle 40 can be used to compress a body part to minimize the thickness through which x-rays pass and to help reduce movement of the body part due to patient movement. X-ray system 10 can also include an object support (not shown) on which the body part can be positioned.

[0022] Digital mammography system 100 can also include a workstation 43, as is known in the art, to receive and process the image data from detector 18. Workstation 43 can include a processor and a memory, and can be configured to perform various image processing operations on the image data, such as image reconstruction, image enhancement, image segmentation, and the like. Workstation 43 can also include a display to display the images to a user. Figure 2Further illustrated, workstation 43 may include a controller 44, which includes at least one processor and memory. The controller 44 may be communicatively coupled to one or more components of the x-ray system 10, including one or more of the radiation source 16, radiation detector 18, pressure paddle 40, and biopsy equipment. In some examples, communication between the controller and the x-ray system 10 may be via a wireless communication system. In other examples, the controller 44 may be electrically communicated with one or more components of the x-ray system via cable 47. Furthermore, in exemplary embodiments, such as Figure 2 As shown, controller 44 is integrated into workstation 43. In other embodiments, controller 44 may be integrated into one or more of the various components of the system 10 disclosed above. Additionally, controller 44 may include processing circuitry that executes stored program logic and may be any of various computers, processors, controllers, or combinations thereof that are compatible with and usable with the various types of equipment and devices used in the x-ray system 10.

[0023] Workstation 43 may include a radiation shield 48 that protects the operator of system 10 from radiation emitted by radiation source 16. Workstation 43 may also include a display 50, a keyboard 52, a mouse 54 and / or other suitable user input devices that facilitate the control of system 10 via user interface 56.

[0024] The controller 44 can adjust the operation and functions of the x-ray system 10. For example, the controller 44 can provide timing control on when the x-ray source 16 emits x-rays, and can further adjust how the detector 18 reads and transmits information or signals after the x-rays hit the detector 18, and how the x-ray source 16 and the detector 18 move relative to each other and relative to the body part being imaged. The controller 44 can also control how information (including images 42 and data acquired during operation) is processed, displayed, stored, and manipulated. The actions performed by the controller 44, as described herein, are related to... Figures 5 to 6 The various method steps described herein can be provided by a set of instructions stored in the non-transitory memory of the controller 44.

[0025] Further, as described above, radiation detector 18 receives radiation rays emitted by radiation source 16. Specifically, during imaging with an X-ray system, a projected image of the body part being imaged can be obtained at detector 18. In some embodiments, data received by radiation detector 18, such as projected image data, can be transmitted from radiation detector 18 to controller 44 via electrical and / or wireless communication. Controller 44 can then reconstruct one or more scanned images based on the projected image data, for example, by implementing a reconstruction algorithm. The reconstructed image can be displayed to a user on user interface 50 via display screen 56.

[0026] The radiation source 16, together with the radiation detector 18, forms part of an x-ray system 10 that provides x-ray images for one or more of screening abnormalities, diagnosis, dynamic imaging, and image-guided biopsy. For example, the x-ray system 10 can operate in a mammography mode to screen for abnormalities. During mammography, a patient’s breast is positioned and compressed between the detector 18 and the compression paddle 40. Thus, the volume of the x-ray system 10 between the compression paddle 40 and the detector 18 is the imaging volume. The radiation source 16 then emits radiation rays onto the compressed breast, and a projection image of the breast is formed on the detector 18. The projection image can then be reconstructed by the controller 44 and displayed on the interface 50. During mammography, the gantry 15 can be adjusted at different angles to obtain images of different orientations, such as cranio-caudal (CC) images and mediolateral oblique (MLO) images. In one example, the gantry 15 can be rotated about the axis 58 while the compression paddle 40 and the detector 18 remain stationary. In other examples, the gantry 15, the compression paddle 40, and the detector 18 can be rotated as a single unit about the axis 58.

[0027] Additionally, the x-ray system 10 can operate in a tomosynthesis mode for performing digital breast tomosynthesis (DBT). During tomosynthesis, the x-ray system 10 can be operated to direct low-dose radiation at various angles within an angular range of the x-ray system 10 to the imaging volume (between the compression paddle 40 and the detector 18). In particular, during tomosynthesis, similar to mammography, the breast is compressed between the compression paddle 40 and the detector 18. The radiation source 16 is then rotated from -0 to +0, and a plurality of projection images of the compressed breast are obtained at regular angular intervals over the angular range. For example, if the angular range of the x-ray system is ±11 degrees, then during an angular sweep of the gantry, approximately every degree swept, the detector can capture 22 projection images, generating a set of angularly x-ray images. The plurality of projection images are then processed by the controller 44 to generate a plurality of DBT image slices. This processing can include applying one or more reconstruction algorithms to reconstruct a three-dimensional image of the breast. Furthermore, the x-ray system can be configured to perform a DBT-guided biopsy procedure. Thus, in some example embodiments, the system 10 can also include a biopsy device including a biopsy needle for extracting a tissue sample for further analysis.

[0028] In some examples, the digital mammography system 100 can be configured to perform contrast imaging, in which a contrast agent, such as iodine, can be injected into the patient and reach a region of interest (ROI) within the breast (e.g., a lesion). The contrast agent is absorbed in blood vessels surrounding a cancerous lesion in the ROI, providing a contrast image for a period of time relative to the surrounding tissue, thereby enhancing the ability to locate the lesion.

[0029] In some embodiments, the digital mammography system 100 includes or is coupled to a picture archiving and communication system (PACS). In one example implementation, the PACS is further coupled to a remote system such as a radiology information system, a hospital information system, and / or to an internal or external network (not shown) to allow operators in different locations to provide commands and parameters and / or to gain access to image data.

[0030] In some examples, the images reconstructed by the controller 44 can store the reconstructed images in a storage device. Alternatively, the images can be transmitted to generate useful patient information for diagnosis and evaluation. In certain embodiments, the controller 44 can transmit the reconstructed images and / or patient information to a display 50. In some embodiments, the reconstructed images can be transmitted from the controller 44 to a storage device for short-term or long-term storage.

[0031] The various methods and processes described further herein, such as the methods described below with reference to Figure 5 and Figure 6 may be stored as executable instructions in a non-transitory memory on a computing device (or controller) in an imaging system (e.g., the digital mammography system 100). In one embodiment, the controller 44 can include such executable instructions in the non-transitory memory and can apply the methods described herein to reconstruct images from scan data. Further, the methods described herein for reading out pixels and merging images can be stored and executed by the controller 44.

[0032] It should be understood that while digital mammography systems are described herein, other x-ray systems using flat panel detectors, such as CT imaging systems, x-ray systems, fluoroscopy systems, interventional radiology systems, etc., can also be applied without departing from the scope of the present disclosure. Figure 1 and Figure 2 It should be understood that while digital mammography systems are described herein, other x-ray systems using flat panel detectors, such as CT imaging systems, x-ray systems, fluoroscopy systems, interventional radiology systems, etc., can also be applied without departing from the scope of the present disclosure.

[0033] Turning now to Figure 3 , a schematic diagram of a photodiode pixel pair 300 of an x-ray system including a detector is shown. The photodiode pixel pair 300 can be a photodiode pixel array, such as the photodiode pixel array 200 of the digital mammography system 100. The photodiode pixel pair 300 includes a first photodiode pixel 302 and a second photodiode pixel 304. The first photodiode pixel 302 and the second photodiode pixel 304 can be arranged in a row and column configuration, such as the row and column configuration of the photodiode pixel array 200. Figures 1 to 2one of a plurality of pairs of photosensor pixels in a photosensor pixel array of the detector 18. As noted, the photosensor array can include a plurality of photosensor pixels, each of which can be a rectangular pixel. The pixels can be arranged in pairs of adjacent pixels. In some examples, the pairs can be predefined and known to the system. In other examples, the pairs can be flexible so as to vary which pixels comprise a pair depending on the application and use. The pixels in the array can be arranged in rows and columns, typically in an N x M matrix arrangement, with each photosensor connected to an address line and a scan line to allow for the generation of a charge in each respective photosensor in response to incident light to be read.

[0034] Each pixel in the photosensor pixel array can include a thin film transistor (TFT) (e.g., a field effect transistor), a TFT control line, a data readout line, a common voltage line, and one or more photodiodes (e.g., a photosensitive p-i-n photodiode). For example, the pair of photosensor pixels 300 can include a first pixel 302 and a second pixel 304. The first pixel can include a first TFT 312, a first TFT control line 320, a first data readout line 324, a first common voltage line 328, and a first photodiode 316. The second pixel 304 can include a second TFT 314, a second TFT control line 322, a second data readout line 324, a second common voltage line 330, and a second photodiode 318. The first TFT 312 and the second TFT 314 can act as switches to individually “turn on” or “turn off” each pixel. The first TFT control line 320 and the second TFT control line 322 can be energized or de-energized to turn on or turn off the respective TFTs, which can allow the photodiodes to be read out one at a time or simultaneously depending on the application. The components of each of the pixels as described herein can allow each pixel to generate a respective image data signal.

[0035] Each of the first pixel 302 and the second pixel 304 can have a first dimension 310. In some examples, the first dimension 310 can be a pixel height or a vertical dimension, as shown. However, in other examples, the first dimension 310 can be a width or a horizontal dimension. The first pixel 302 can have a second dimension 306 that is perpendicular to the first dimension 310. The second pixel 304 can also have a corresponding second dimension 308 that is perpendicular to the first dimension 310. The second dimension 306 of the first pixel 302 and the second dimension 308 of the second pixel 304 can be equal in length, and either can be considered the second dimension of any pixel of the detector array. Figure 3

[0036] ​In some examples, the second dimension 306, 308 can be half the length of the first dimension 310. As an example, the first dimension 310 can be 100 pm, and the second dimension 306, 308 can each be 50 pm, such that each of the first pixel 302 and the second pixel 304 is 50 pm by 100 pm. Thus, the pixels can be rectangular pixels, and the pixel pair can be square or approximately square. It will be appreciated that other dimensions are possible without departing from the scope of the present disclosure.

[0037] When arranged in a pixel matrix, the first pixel 302 and the second pixel 304 in the photo sensor pixel pair 300 can be directly adjacent to one another. As will be described with reference to Figure 5 and Figure 6 Depending on whether the application selected is a low dose application or a high dose application, the first TFT control line 320 and the second TFT control line 322 can be energized simultaneously or sequentially. For example, in a low dose application, the control lines can be energized simultaneously, and thus the photo sensor pixel pair 300 can be read out together by the data readout lines 324, 326 to form one image having a total pixel size of the first dimension 310 times the second dimension 306 plus the second dimension 308 (e.g., for a total pixel size of 100 pm by 100 pm). In a high dose application, the control lines can be energized sequentially, and thus the first pixel 302 and the second pixel 304 can be read out separately by the data readout lines 324, 326. Additionally, as will be explained further below with respect to Figure 4 in a high dose application, the detector can be translated by half the first dimension 310 (e.g., 50 pm when the first dimension 310 is 100 pm). Energizing the two control lines in the pair sequentially and translating the detector can thus generate two images of the pixels with an overlap of 50 pm.

[0038] For rectangular pixels having a pitch of 50 pm by 100 pm, the area of the photodiode can be a-Si compatible. For example, in some examples, the minimum area of a pixel that can be used with a-Si photodiodes can be approximately 5000 pm 2 . Rectangular pixels having a pitch of 50 pm by 100 pm as presented herein can have an area of 5000 pm 2 which can be a-Si photodiode compatible. Thus, rectangular pixels can allow for high image quality to be achieved in both low dose applications and high dose applications, while allowing for the use of readily available a-Si photodiodes during manufacturing.

[0039] Turning now to Figure 4 , a pixel matrix 400 is shown with respect to Figure 3a schematic diagram 400 of a detector position of a photodiode pixel pair 300. The photodiode pixel pair 300 can include a first pixel 302 and a second pixel 304, as previously described. The first pixel 302 and the second pixel 304 can be symmetric rectangular pixels having dimensions as previously described. As a non-limiting example, the pixels can have dimensions of 100 pm x 50 pm.

[0040] The schematic diagram 400 shows a first position 402 of the detector and a second position 404 of the detector in a high dose application. As previously described, the photodiode pixel pair 300 is representative of a plurality of pixel pairs that make up a pixel array. The dashed lines of the first position 402 and the second position 404 of the detector can represent a spatial extent of the pixel pair that is imaged by the system with the detector in a given position, and are not intended to depict a shape or actual dimensions of the detector.

[0041] As described above, the pixel array can include a matrix of pixel pairs. The first position 402 of the detector can image the matrix of pixel pairs in the first position. The second position 404 can be a translation of the detector by half the length of the first dimension 310 of the pixel pair 300. For example, the second position 404 can be a first distance 406 translated from the first position 402, the first distance 406 being half of the first dimension 310. As an example, when the first dimension 310 is 100 pm, the first distance 406 can be 50 pm.

[0042] As will be further described below, in a high dose application, two images are generated and the detector is translated by the first distance 406 between the two images. Thus, with each pixel in the pixel pair 300 read out individually for each image, two images can be generated having rectangular pixel dimensions. The two images can have overlapping data and can be merged together to form one image having pixels with a dimension of half the first dimension 310 multiplied by a second dimension (e.g., the second dimensions 306 and / or 308) using one or more image processing techniques. For example, for rectangular pixels of 100 pm x 50 pm, generating two images and translating the detector by 50 pm along the 100 pm dimension can result in an effective pixel pitch of 50 pm x 50 pm.

[0043] Turning now to Figure 5 a flowchart illustrating a method 500 for generating an x-ray based image of a patient for a low dose application. The method 500 can use the system described above with respect to the system 100. The method 500 can include a first image 502, a second image 504, and a third image 506. The first image 502 can be generated by the system 100 using the first position 402 of the detector. The second image 504 can be generated by the system 100 using the second position 404 of the detector. The third image 506 can be generated by the system 100 using the first position 402 of the detector. Figures 1 to 3The system and components described herein are used to perform the method; however, it should be understood that similar methods can be used with other systems without departing from the scope of this disclosure. Method 500 can be executed via instructions stored in non-transitory memory of one or more computing devices. For example, instructions can be stored in memory and executed by… Figure 1 The controller 44 is executed by one or more processors. Method 500 is described relative to rectangular pixels having dimensions of 50 μm × 100 μm; however, it should be understood that other rectangular pixel dimensions may be used without departing from the scope of this disclosure.

[0044] At 502, method 500 includes receiving a request to initiate a patient scan. This request may be made by a user via communication with an imaging system (e.g., Figure 1 The user equipment (e.g., via the x-ray system 10) communicates with the user equipment. Figure 1 The imaging system may include a detector with a photoelectric sensor pixel array comprising a rectangular pixel matrix divided into multiple pixel pairs, such as relative to... (the operator console of workstation 43) user selection or other user input. Figure 3 Described. Each of these rectangular pixels may include a TFT and a TFT control line.

[0045] At 504, method 500 includes determining the requested imaging system scheme. For example, the request may indicate the scan type, scan scheme, and other parameters that may indicate to the imaging system the requested view, the requested irradiance, etc. As a non-limiting example, scan schemes for mammography applications may include 2D mammography schemes, 3D mammography schemes, and 3D tomographic fusion schemes, each of which may require different imaging settings, irradiance, etc.

[0046] At 506, method 500 determines whether a low-dose protocol is requested. As an example, the requested imaging system protocol, as determined at 504, may indicate the radiation dose to be delivered to the patient during image acquisition. In some examples, the imaging system may include a threshold radiation dose, wherein the protocol may be low-dose if the requested protocol requires a dose below the threshold, and high-dose if the requested protocol requires a dose above the threshold. In other examples, each imaging protocol known to the system may be designated as low-dose or high-dose. For example, screening protocols, such as screening mammograms or screening chest CT scans, may be predefined as low-dose. If the requested protocol is low-dose, method 500 proceeds to 508. If the requested protocol is not low-dose (e.g., high-dose), method 500 proceeds to 514 to perform methods for high-dose applications, such as relative to... Figure 6 Further description.

[0047] At 508, the method 500 includes energizing the TFT control lines of each pixel pair simultaneously. As noted above, the photosensor pixel array of the detector can include multiple pairs of rectangular pixels, each of the pixels in each pair of pixels including a TFT and a TCT control line configured to turn the pixel on and off. The TFT control lines of each pixel pair can be energized simultaneously such that the pair of pixels can be considered one pixel.

[0048] At 510, the method 500 includes reading out each pixel pair together. Reading out the pixel pair can generate an imaging data signal for each pixel pair. Each of the rectangular pixels can be 50 pm x 100 pm. Reading out the pixel pair together can result in an effective pixel pitch of 100 pm x 100 pm. Thus, the effective pixel pitch in low dose applications can be twice the rectangular pixel pitch.

[0049] At 512, the method 500 includes generating an image having an effective pixel pitch resulting from reading out each pixel pair together. With an effective pixel pitch of 100 pm x 100 pm, which can be a relatively large pixel pitch, low dose acquisition can have a high signal-to-noise ratio and improved image quality.

[0050] Turning now to Figure 6 , a flowchart illustrating a method 600 for generating an x-ray based image of a patient is shown for high dose applications. The method 600 can be performed using the systems and components described herein above with respect to Figures 1 to 3 , it should be understood that similar methods can be used with other systems without departing from the scope of the present disclosure. The method 600 can be performed via instructions stored in a non-transitory memory of one or more computing devices. For example, the instructions can be stored in memory and executed by one or more processors of the controller 44 of the Figure 1 .

[0051] At 602, the method 600 includes receiving a request to initiate a scan of a patient. The request can be a user selection or other user input entered by a user via a user device in communication with an imaging system (e.g., the x-ray system 10 of Figure 1 ) such as via an operator console of the workstation 43 of Figure 1 . The imaging system can include a detector having a photosensor pixel array including a matrix of rectangular pixels divided into multiple pairs of pixels as described with respect to Figure 3 . Each of the rectangular pixels can include a TFT and a TFT control line.

[0052] At 604, the method 600 includes determining a requested imaging system protocol. For example, the request can indicate a scan type, a scan protocol, and other parameters that can indicate a requested view, a requested radiation dose, etc. to the imaging system. As a non-limiting example, scan protocols for mammography applications can include a 2D mammogram image protocol, a 3D mammogram image protocol, and a 3D tomosynthesis protocol, each of which can require different imaging settings, radiation dose, etc.

[0053] At 606, the method 600 determines whether a high dose protocol is requested. As an example, the requested imaging system protocol as determined at 604 can indicate a radiation dose to be delivered to a patient when acquiring images. In some examples, the imaging system can include a threshold radiation amount, where if the requested protocol requires a radiation dose below the threshold, the protocol can be a low dose, and if the requested protocol requires a radiation dose above the threshold, the protocol can be a high dose. In other examples, each imaging protocol known to the system can be designated as a low dose or a high dose. For example, diagnostic exam protocols, such as diagnostic mammogram images, can be predefined as low dose. If the requested protocol is a low dose, the method 600 proceeds to 608. If the requested protocol is not a low dose (e.g., a high dose), the method 600 proceeds to 618 to perform a method for low dose applications, as will be described. Figure 5

[0054] At 608, the method 600 includes energizing the first TFT control line and the second TFT control line of each pixel pair for a first x-ray. As described above, energizing the TFT control lines in sequence can include energizing the first TFT control line at a first time, and energizing the second TFT control line at a second time. The energizing for the first x-ray at 608 can be the energizing at the first time.

[0055] At 610, the method 600 includes generating a first x-ray image. The first x-ray image can correspond to a first position of the detector before the detector is translated. The first x-ray image can include data for the pixels when the detector is at the first position. The energized pixels can be rectangular, for example, 50 pm by 100 pm in size.

[0056] At 612, the method 600 includes translating the detector of the imaging system. For 100 pm by 50 pm rectangular pixels, the detector can be translated 50 pm along the 100 pm dimension. For example, the rectangular pixels can have a vertical dimension of 100 pm and a horizontal dimension of 50 pm. In such examples, the detector can be translated 50 pm along the vertical direction. In some examples, the detector can be translated using a fast piezoelectric linear actuator.

[0057] ​In some examples, an x-ray pulse emitted from an x-ray source to be detected by an x-ray detector can be split into two halves. In such examples, when the x-ray system is stopped (e.g., not actively emitting radiation rays), the detector can be translated midway between the two halves of the pulse, resulting in generated images that are not blurred. Alternatively, the detector can be translated during the pulse without splitting the x-ray pulse. For example, the detector can be translated for a duration of 50 milliseconds, resulting in a total blur of 5% for a 1 second acquisition.

[0058] At 614, the method 600 includes energizing the first TFT control line and the second TFT control line for each pixel pair for a second x-ray. As described above, energizing the TFT control lines in sequence can include energizing the first TFT control line at a first time and energizing the second TFT control line at a second time. The energizing at 614 for the second x-ray can be the energizing at the second time. The pixels during the energizing at the second time can also be rectangular pixel dimensions, such as 50 pm x 100 pm.

[0059] At 616, the method 600 includes generating a second x-ray image. The second x-ray image can correspond to a second position of the detector prior to the detector being translated. The second x-ray image can include data for the pixels when the detector is at the second position. The pixels that are energized can be rectangular, such as 50 pm x 100 pm in size.

[0060] Generating the first image and the second image can include individually reading out each pixel in each pixel pair. Since each pixel includes a data readout line, each pixel can be individually read out. Each of the rectangular pixels can be 50 pm x 100 pm. Individually reading out the pixel pairs can result in a pixel pitch that is maintained at 50 pm x 100 pm. Thus, individually reading out the pixels can result in two imaging data signals being generated per pixel pair, one for each of the pixels in the pair. Furthermore, when the detector is translated, as described above, each pixel can be read out twice, and thus each pixel can generate its two image data signals, one for each of the pixels in each image. The first image and the second image can include data for the same pixels, and thus can be considered oversampled images. In examples where the x-ray pulse is split into two, the first image can correspond to a first half of the x-ray pulse, and the second image can correspond to a second half of the x-ray pulse.

[0061] At 618, the method 600 includes merging the two images together to form a single image. The single image can have an effective pixel pitch of 50 pm x 50 pm, which can be half of a rectangular pixel pitch of 50 pm x 100 pm. Merging the two images can be accomplished via one or more image processing techniques, such as inverse filtering (e.g., deblurring) and / or noise regularization (e.g., to reduce noise enhancement when the signal-to-noise ratio is too low).

[0062] As an example, a filter can be applied to the images. The filter can be a finite impulse response filter, where the desired frequency characteristics of the filter are determined from information about the imaging system. The filter can include an inverse filtering portion and a noise regularization portion controlled by a single parameter. The inverse filtering can reverse blurring effects in the image acquisition data, such as blurring effects from the scintillator and the aperture. The blurring can be modeled as a linear shift-invariant process and can be expressed as a convolution of the original image with a blurring function. The regularization portion of the filter can decrease the response of the filter as the frequency increases to prevent noise enhancement in low signal-to-noise ratio areas. It should be understood that the methods for merging and processing images as presented herein are non-limiting examples, and other suitable methods for merging and processing can be used without departing from the scope of the present disclosure.

[0063] As noted, the image resulting from combining the two images can have an effective pixel pitch of 50 pm x 50 pm. Thus, the effective pixel pitch of 50 pm x 50 pm can be half of a rectangular pixel pitch of 50 pm x 100 pm. In this way, the rectangular pixels disclosed herein can provide a smaller pixel size in high dose applications. With a smaller pixel size, such as 50 pm x 50 pm, high dose applications can generate images with high spatial resolution.

[0064] A technical effect of the systems and methods presented herein is that rectangular pixels can be used to generate smaller sized pixels in high dose applications, allowing for high spatial resolution therein, and larger sized pixels in low dose applications, allowing for high signal-to-noise ratio therein. Thus, a single x-ray system incorporating rectangular pixels as described herein can be used for both low dose imaging applications and high dose imaging applications. As an example use case, mammography imaging can be high dose or low dose, depending on the end use. For example, 3D mammography imaging can be a high dose application, while screening images, such as 2D and / or tomosynthesis mammography, can be a low dose application. With the rectangular pixel detector array presented herein, the same imaging system can be used to acquire applications with different doses. Further, the rectangular pixels as described herein can accommodate a-Si photodiodes, which can improve the manufacturing efficiency of imaging systems capable of generating high quality images in high dose applications.

[0065] The present disclosure also provides support for an x-ray imaging system including: an array of photosensor pixels of a flat panel detector, the array of photosensor pixels including a plurality of pixels having a rectangular pixel pitch, wherein the plurality of pixels are arranged in pairs, and each pixel of the plurality of pixels is configured to generate a respective image data signal via a thin film transistor (TFT), a TFT control line, and a data readout line, wherein in a low dose application, the TFT control lines of the pixels of each pixel pair are simultaneously energized to generate the respective image data signal having a first effective pixel pitch that is twice the rectangular pixel pitch, and in a high dose application, the TFT control lines of the pixels of each pixel pair are sequentially energized to generate the respective image data signal having the rectangular pixel pitch; and a controller coupled to the array of photosensor pixels to receive the respective image data signals generated by the plurality of pixels, wherein, in the high dose application, the controller is configured to translate the flat panel detector between a first position and a second position to generate two image data signals for each pixel of the plurality of pixels, and the controller is configured to combine the two image data signals to generate an image data signal having a second effective pixel pitch that is half the rectangular pixel pitch. In a first example of the system, the rectangular pixel pitch is 50 pm x 100 pm, the first effective pixel pitch in the low dose application is 100 pm x 100 pm, and the second effective pixel pitch in the high dose application is 50 pm x 50 pm. In a second example of the system (optionally including the first example), in the high dose application, a first image is generated with the flat panel detector in the first position, and a second image is generated with the detector in the second position. In a third example of the system (optionally including one or both of the first example and the second example), the flat panel detector is configured to translate 50 pm along the 100 pm dimension of the plurality of pixels between the first position and the second position in the high dose application. In a fourth example of the system (optionally including one or more or each of the first example through the third example), the controller is configured with one or more of an inverse filter and a noise regularization filter stored in a memory that, when executed, combine the two image data signals in the high dose application. In a fifth example of the system (optionally including one or more or each of the first example through the fourth example), each pixel of the plurality of pixels includes an amorphous silicon (a-Si) photodiode. In a sixth example of the system (optionally including one or more or each of the first example through the fifth example), the x-ray imaging system is configured as a mammography imaging system, and wherein the low dose application includes a tomosynthesis application, and the high dose application includes a three-dimensional (3D) application.

[0066] The present disclosure also provides support for a method comprising: determining an imaging system regime of a requested scan; in response to determining the imaging system regime, energizing thin film transistor (TFT) control lines of rectangular pixels in each pair of rectangular pixels in a pixel array of a detector in one of a simultaneous and sequential manner, wherein the TFT control lines are energized simultaneously in a low dose regime and sequentially in a high dose regime; generating an imaging data signal, wherein the imaging data signal comprises an imaging data signal of each pair when the TFT control lines of each pair are energized simultaneously and the imaging data signal comprises two imaging data signals of each pair when the TFT control lines of each pair are energized sequentially. In a first example of the method, the rectangular pixels have a pixel pitch of 50 pm x 100 pm. In a second example of the method (optionally comprising the first example), generating the imaging data signal in the low dose regime comprises reading out pairs of pixels together for an effective pixel pitch of 100 pm x 100 pm. In a third example of the method (optionally comprising one or both of the first example and the second example), generating the imaging data signal in the low dose regime comprises reading out the rectangular pixels in each pair individually with the pixel pitch of 50 pm x 100 pm. In a fourth example of the method (optionally comprising one or more or each of the first example through the third example), the method further comprises, when the imaging system regime is the high dose regime: translating the detector by 50 pm along a first direction; generating two images from the rectangular pixels read out with the pixel pitch of 50 pm x 100 pm; and combining the two images into one image, wherein the one image has an effective pixel pitch of 50 pm x 50 pm. In a fifth example of the method (optionally comprising one or more or each of the first example through the fourth example), combining the two images into one image comprises applying one or more of an inverse filter and a noise regularization filter.

[0067] The present disclosure also provides support for an imaging system comprising: an x-ray detector comprising a matrix of rectangular pairs of pixels, wherein each pixel in each pixel pair comprises a thin film transistor (TFT), a TFT control line, and a data readout line, wherein each pixel in each rectangular pair of pixels is configured with a first dimension and a second dimension that is half the length of the first dimension. In a first example of the system, each pixel in each rectangular pair of pixels has a pixel pitch of 50 pm x 100 pm, wherein the first dimension is 100 pm and the second dimension is 50 pm. In a second example of the system, optionally including the first example, the TFT control lines of each pixel in each rectangular pair of pixels are simultaneously energized in a low dose application. In a third example of the system, optionally including one or both of the first and second examples, the TFT control lines of each pixel in each rectangular pair of pixels are sequentially energized in a high dose application. In a fourth example of the system, optionally including one or more or each of the first through third examples, the x-ray detector is configured to translate half the length of the first dimension in a direction along the first dimension. In a fifth example of the system, optionally including one or more or each of the first through fourth examples, the x-ray detector is configured to receive a pulse of radiation rays of emitted radiation from a radiation source that emits the radiation rays, wherein in a high dose application, the radiation source is configured to separate the pulse of the radiation rays into two halves, and wherein the x-ray detector is configured to translate half the length of the first dimension in a direction along the first dimension between the two halves of the pulse. In a sixth example of the system, optionally including one or more or each of the first through fifth examples, the x-ray detector is configured to receive a pulse of radiation rays of emitted radiation from a radiation source that emits the radiation rays, wherein in a high dose application, the x-ray detector is configured to translate the half the length of the first dimension in a direction along the first dimension during the pulse of the radiation rays.

[0068] As used herein, an element or step recited in the singular and preceded with the word "a" or "an" should be understood as not excluding plural of said elements or steps, unless explicitly stated that such exclusion applies. Also, aspects of the disclosure can include "one embodiment," or "an embodiment," or "some embodiments," and the phrases "comprises," "comprising," "includes," "including," or "has," along with their derivatives, can indicate that additional such elements or steps are included. The terms "including" and "in which" are used as the plain language equivalents of the respective terms "comprising" and "wherein." Also, the terms "first," "second," or "third," etc. are used merely as labels, and are not intended to impose numerical requirements on their objects.

[0069] This written description uses examples to disclose the application, including the best mode, and also to enable any person skilled in the art to practice the application, including making and using any devices or systems and performing any incorporated methods. The patentable scope of the application is defined by the claims, and can include other examples that occur to those skilled in the art. Such other examples are intended to fall within the scope of the claims if they have structural elements that do not differ from the literal language of the claims, or if they include equivalent

Claims

1. An x-ray imaging system (100), comprising: a photosensor pixel array of a flat panel detector (18), the photosensor pixel array comprising a plurality of pixels having a rectangular pixel pitch, wherein the plurality of pixels are arranged in pairs (300) and each pixel of the plurality of pixels is configured to generate a respective image data signal via a thin film transistor (TFT) (312), a TFT control line (320), and a data readout line (324), wherein in a low dose application, the TFT control lines of the pixels in each pixel pair are simultaneously energized to generate the respective image data signal having a first effective pixel pitch that is twice the rectangular pixel pitch, and in a high dose application, the TFT control lines of the pixels in each pixel pair are sequentially energized to generate the respective image data signal having the rectangular pixel pitch; and a controller (44) coupled to the photosensor pixel array to receive the respective image data signals generated by the plurality of pixels, wherein, in the high dose application, the controller is configured to translate the flat panel detector (18) between a first position (402) and a second position (404) to generate two image data signals for each pixel of the plurality of pixels, and the controller is configured to combine the two image data signals to generate an image data signal having a second effective pixel pitch that is half the rectangular pixel pitch.

2. The x-ray imaging system of claim 1, wherein the rectangular pixel pitch is 50 pm x 100 pm, the first effective pixel pitch in a low dose application is 100 pm x 100 pm, and the second effective pixel pitch in a high dose application is 50 pm x 50 pm.

3. The x-ray imaging system of claim 1, wherein, In a high dose application, a first image is generated with the flat panel detector (18) in the first position (402) and a second image is generated with the detector (18) in the second position (404).

4. The x-ray imaging system of claim 1, wherein the flat panel detector (18) is configured to translate 50 pm between the first and second positions (402, 404) along the 100 pm dimension of the plurality of pixels (300) in a high dose application.

5. The x-ray imaging system of claim 1, wherein the controller (44) is configured with one or more of an inverse filter and a noise regularization filter stored in a memory that, when executed, combine the two image data signals in a high dose application.

6. The x-ray imaging system of claim 1, wherein each pixel of the plurality of pixels comprises an amorphous silicon (a-Si) photodiode (316).

7. The x-ray imaging system of claim 1, wherein the x-ray imaging system (100) is configured as a mammography imaging system, and wherein the low dose application comprises a tomo-fusion application and the high dose application comprises a three-dimensional (3D) application.

8. A method comprising: determining an imaging system regime of a requested scan (504); in response to determining the imaging system regime, energizing thin film transistor (TFT) control lines of rectangular pixels in each pair of rectangular pixels in a pixel array of a detector in one of a simultaneous and sequential manner, wherein TFT control lines are energized simultaneously in a low dose regime (508) and sequentially in a high dose regime (608, 614); generating an imaging data signal, wherein the imaging data signal comprises an imaging data signal for each pair when the TFT control lines of each pair are energized simultaneously and the imaging data signal comprises two imaging data signals for each pair when the TFT control lines of each pair are energized sequentially.

9. The method of claim 8, wherein the rectangular pixels (302, 304) have a pixel pitch of 50 pm x 100 pm.

10. The method of claim 9, wherein generating an imaging data signal in a low dose regime comprises reading out pairs of pixels together for an effective pixel pitch of 100 pm x 100 pm.

11. The method of claim 9, wherein generating an imaging data signal in a low dose regime comprises reading out rectangular pixels in each pair individually with the pixel pitch of 50 pm x 100 pm.

12. The method of claim 11, further comprising, when the imaging system regime is a high dose regime: translating the detector along a first direction by 50 pm (612); generating two images from the rectangular pixels read out with the pixel pitch of 50 pm x 100 pm (610, 616); and combining the two images into one image, wherein the one image has an effective pixel pitch of 50 pm x 50 pm (618).

13. The method of claim 12, wherein combining the two images into one image comprises applying one or more of an inverse filter and a noise regularization filter (618).

14. The method of claim 9, further comprising: receiving the radiation rays from a radiation source that emits pulses of the radiation rays (608, 616), wherein in a high dose application, the radiation source is configured to separate the pulses of the radiation rays into two halves, and wherein the detector is configured to translate along a direction of the 100 pm dimension by half a length of the 100 pm dimension between the two halves of the pulses (612).

15. The method of claim 9, further comprising: receiving radiation from a radiation source that emits pulses of the radiation rays (608, 616), wherein in a high dose application, the detector is configured to translate along a direction of the 100 pm dimension by half a length of the 100 pm dimension during the pulses of the radiation rays (612).

14. The method of claim 13, wherein the radiation source is a computed tomography (CT) scanner.

15. The method of claim 13, wherein the radiation source is a digital breast tomosynthesis (DBT) scanner.

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