Reduction of image lag in X-ray detector panels

Abstract

A radiation therapy system is configured with a fast readout of an X-ray image with significantly reduced image lag. A reset phase is included in the process of acquiring an X-ray image to reduce image lag in subsequently acquired X-ray images. During the reset phase, residual charge is concurrently transferred from a plurality of pixel detector element arrays of an X-ray detector panel. As a result, image lag present in subsequent X-ray images is minimized or otherwise reduced.

Description

Background Art

[0001] Unless otherwise indicated herein, the approaches described in this section are not prior art to the claims in this application and are not admitted to be prior art by inclusion in this section.

[0002] Radiation therapy is a localized treatment directed at a specific target tissue (planned target volume), such as a cancerous tumor. Ideally, radiation therapy is performed on the planned target volume so that the surrounding normal tissue is protected from receiving a dose above a specified tolerance, thereby minimizing the risk of damage to healthy tissue. Prior to the implementation of radiation therapy, an imaging system is typically employed to provide a three-dimensional image of the target tissue and surrounding area. Based on this imaging, the size and mass of the target tissue can be estimated, and an appropriate treatment plan can be generated and the planned target volume determined.

[0003] In order to correctly deliver the prescribed dose to the planned target volume (i.e., target tissue) during radiotherapy, the patient should be correctly positioned relative to the linear accelerator that delivers the radiotherapy. Typically, dosimetry and geometric data are checked before and during treatment to ensure correct patient placement and that the radiotherapy treatment delivered matches the previously planned treatment. This process is called image-guided radiation therapy (IGRT) and involves using an imaging system to view the target tissue during or before delivering the radiotherapy to the planned target volume. IGRT incorporates imaging coordinates from the treatment plan to ensure that the patient is correctly aligned for treatment in the radiotherapy device.

[0004] Public Content

[0005] According to at least some embodiments of the present disclosure, a radiation therapy system is configured with a fast readout of an X-ray image without significantly increasing image lag. In an embodiment, a reset phase is included in the process of acquiring an X-ray image to reduce image lag in subsequently acquired X-ray images. Specifically, during the reset phase, residual charge is concurrently transferred from a plurality of pixel detector element arrays in an X-ray detector panel. As a result, image lag present in subsequent X-ray images is minimized or otherwise reduced.

[0006] The foregoing summary is illustrative only and is not intended to be limiting in any way. In addition to the illustrative aspects, embodiments, and features described above, further aspects, embodiments, and features will become apparent by reference to the drawings and the following detailed description. BRIEF DESCRIPTION OF THE DRAWINGS

[0007] The foregoing features and other features of the present disclosure will become more apparent from the following description and the appended claims in conjunction with the accompanying drawings. These drawings depict only a few embodiments according to the present disclosure and therefore should not be considered as limiting the scope thereof. The present disclosure will be described with additional specificity and detail through the use of the accompanying drawings.

[0008] Figure 1 is a perspective view of a radiation therapy system in which various aspects of the present disclosure may be beneficially implemented.

[0009] Figure 2 Schematically illustrating various embodiments according to the current disclosure Figure 1 The drive stand and gantry of the radiation system.

[0010] Figure 3 A cross-sectional view of an X-ray imager according to an embodiment of the present disclosure is schematically shown.

[0011] Figure 4 Schematically illustrating various embodiments disclosed herein based on Figure 1 The RT system includes one or more X-ray imagers that generate projection images to construct the digital volume.

[0012] Figure 5 According to one embodiment of the present disclosure Figure 3 A partial circuit diagram of the photosensitive element and detector circuit layer included in the X-ray imager.

[0013] Figure 6 According to another embodiment of the present disclosure Figure 3 A partial circuit diagram of the photosensitive element and detector circuit layer included in the X-ray imager.

[0014] Figure 7 According to another embodiment of the present disclosure Figure 3 A partial circuit diagram of the photosensitive element and detector circuit layer included in the X-ray imager.

[0015] Figure 8 is a timing diagram schematically illustrating charge accumulation and loss in a pixel detector element during the illumination phase, readout phase, and reset phase of a single X-ray image acquisition according to an embodiment of the present disclosure.

[0016] Fig. 9 The present invention is shown in the embodiment of the present invention. Figure 8 Schematic timing diagram of the charge loss of two adjacent rows of pixel detector elements during the readout phase of .

[0017] Fig.10 The present invention is shown in the embodiment of the present invention. Figure 8Schematic timing diagram of charge loss of two adjacent rows of pixel detector elements during the reset phase of FIG.

[0018] Fig.11 A flow chart of a method for acquiring X-ray image data in an X-ray detector panel according to one or more embodiments of the present disclosure is set forth.

[0019] Fig. 12A is a timing diagram schematically illustrating charge loss of different rows of pixel detector elements during a reset phase according to an embodiment of the present disclosure.

[0020] Fig. 12B FIG. 1 shows a first pixel detector element array and an adjacent second pixel detector element array according to an embodiment of the present disclosure. Fig. 12A part of the reset phase.

[0021] Fig.13 is a timing diagram schematically illustrating the relative timing of the irradiation phase, the readout phase, and the reset phase with respect to the delivery of the therapy beam pulse 1350 according to an embodiment of the present disclosure. DETAILED DESCRIPTION

[0022] In the following detailed description, reference is made to the accompanying drawings which form a part of this description. In the accompanying drawings, similar symbols generally identify similar components unless the context indicates otherwise. The illustrative embodiments described in the detailed description, drawings, and claims are not meant to be limiting. Other embodiments may be utilized, and other changes may be made, without departing from the spirit or scope of the subject matter presented herein. It will be readily understood that the various aspects of the present disclosure as generally described herein and shown in the accompanying drawings may be arranged, substituted, combined, and designed in a variety of different configurations, all of which are expressly intended and form a part of the present disclosure.

[0023] introduction

[0024] Image guided radiation therapy (IGRT) is used to treat tumors in areas of the body subject to voluntary movement (such as the lungs), or in areas of the body subject to involuntary movement (such as organs affected by peristalsis). IGRT involves using an imaging system to view the target tissue (also called a "target volume") while applying radiation therapy to it. In IGRT, the image-based coordinates of the target volume from a previously determined treatment plan are compared with the image-based coordinates of the target volume determined during the application of the treatment beam. In this way, changes in the surrounding organs at risk and / or movement or deformation of the target volume relative to the radiation therapy system can be detected. Therefore, dose limitations on organs at risk are accurately implemented based on the position and shape of each day, and the position of the patient and / or the treatment beam can be adjusted to more accurately target the radiation dose to the tumor. For example, in pancreatic tumor treatment, organs at risk include the duodenum and stomach. The shape and relative position of these organs at risk relative to the target volume may change significantly from day to day. Therefore, accurate adaptation to the shape and relative position of such organs at risk enables dose saving to those organs at risk, increasing the dose to the target volume, and achieving better treatment results.

[0025] In some conventional IGRT radiation systems, soft tissue motion is detected during application of the treatment beam via fiducial markers, such as gold seeds. However, the use of fiducial markers has many disadvantages, particularly the invasive surgical procedure required to place the markers. Specifically, laparoscopic insertion of fiducial markers requires additional time and clinical resources, such as an operating room, anesthesia, antibiotics, and the involvement of many additional medical specialists.

[0026] Alternatively, in some conventional IGRT radiation systems, the movement of soft tissue is detected during the application of the treatment beam via magnetic resonance imaging (MRI). However, MRI-based IGRT also has disadvantages. First, MRI-based IGRT systems are generally larger, more complex and more expensive than radiotherapy systems that use X-ray imaging. Second, detecting the movement or deformation of the target area via MRI generally involves monitoring images associated with 2D slices passing through the target area. As a result, it is difficult to detect target area movement or deformation that occurs anywhere outside the monitored 2D slice (or perpendicular to the 2D slice), which can significantly affect the accuracy of the applied radiation dose.

[0027] Alternatively, in some conventional IGRT radiation systems, the movement of soft tissue is detected during the application of therapeutic X-rays via imaging X-rays that also pass directly through the target area. For example, volumetric image data of the target area can be reconstructed based on X-ray projection images of the target area generated by a computed tomography (CT) or cone beam CT (CBCT) process. During the CT or CBCT process, multiple X-ray projection images are generated by imaging X-rays passing through the target area and arriving at an X-ray detector panel or other X-ray imaging device. Generally, in IGRT applications, faster CT or CBCT acquisition is beneficial because faster acquisition of target area images enables faster detection of movement or deformation of the target area and / or changes in surrounding organs at risk.

[0028] The speed of CT or CBCT acquisition is strongly dependent on the panel readout time of the X-ray detector panel that generates the X-ray projection images of the target volume and surrounding organs at risk. Because all pixels in a row of the X-ray detector panel are typically read out simultaneously, the minimum panel readout time is approximately equal to the pixel readout time multiplied by the number of rows in the X-ray detector panel. Therefore, applying a shorter pixel readout time in the X-ray detector panel can significantly reduce the panel readout time. However, the shorter pixel readout time in some X-ray detector panels (such as amorphous silicon-based panels) will inevitably cause a large image lag in the X-ray images generated by such X-ray detector panels. Image lag refers to the carryover of charge associated with the pixels from one X-ray image to the next X-ray image, and may cause significant image artifacts. Therefore, there is a well-known tradeoff between imaging frame rate and image lag when achieving faster X-ray image acquisition.

[0029] In view of the foregoing, there is a need in the art for improved systems and techniques to increase imaging frame rates in radiation therapy systems without increasing image lag in the resulting X-ray images. Figure 1 One such embodiment is shown in .

[0030] System Overview

[0031] Figure 11 is a perspective view of a radiation therapy system 100 in which various aspects of the present disclosure may be beneficially implemented. The radiation therapy (RT) system 100 is a radiation system configured to detect intra-fraction motion in near real time using X-ray imaging techniques. Thus, the RT system 100 is configured to provide stereotactic radiosurgery and precision radiation therapy for lesions, tumors, and conditions anywhere in the body where radiation therapy is indicated. Thus, the RT system 100 may include one or more linear accelerators (LINACs), kilovolt (kV) X-ray sources, X-ray imagers, and in some embodiments, megavolt (MV) electron portal imaging devices (EPIDs), and one or more linear accelerators (LINACs) generate megavolt (MV) treatment beams of high-energy X-rays. For example, the radiation therapy system 100 is described herein as being configured with a circular gantry. In other embodiments, the radiation therapy system 100 may be configured with a C-shaped gantry capable of infinite rotation via a slip ring connection, a ring gantry with a slip ring, a C-shaped gantry with a winding configuration, and the like.

[0032] Typically, the RT system 100 is capable of kV imaging of the target area during the application of the MV treatment beam, so that the IGRT process can be performed using X-ray imaging rather than MRI. The RT system 100 may include one or more touch screens 101, a couch motion control 102, a bore 103, a base positioning assembly 105, a couch 107 arranged on the base positioning assembly 105, and an image acquisition and treatment control computer 106, all of which are arranged in the treatment room. The RT system 100 also includes a remote control console 110, which is arranged outside the treatment room and enables treatment delivery and patient monitoring from a remote location. The base positioning assembly 105 is configured to accurately position the couch 107 relative to the bore 103, and the motion control 102 includes an input device (such as a button and / or a switch) that enables a user to operate the base positioning assembly 105 to automatically and accurately position the couch 107 to a predetermined position relative to the bore 103. The motion controls 102 also enable a user to manually position the couch 107 to a predetermined position. In certain embodiments, the RT system 100 also includes one or more cameras (not shown) for patient monitoring in the treatment room.

[0033] Figure 2 The drive bracket 200 and the frame 210 of the RT system 100 according to various embodiments of the present disclosure are schematically shown. For clarity, Figure 2The housing, base positioning assembly 105, treatment bed 107 and other components of the RT system 100 are omitted. The drive support 200 is a fixed support structure for the components of the RT treatment system 110, including a gantry 210 and a drive system 201 for rotatably moving the gantry 210. The drive support 200 rests on and / or is fixed to a support surface external to the RT treatment system 110, such as the floor of an RT treatment facility. The gantry 210 is rotationally coupled to the drive support 200 and is a support structure on which various components of the RT system 100 are mounted, including a linear accelerator (LINAC) 204, an MV electronic portal imaging device (EPID) 205, an imaging X-ray source 206, and an X-ray imager 207. During operation of the RT treatment system 110, the gantry 210 rotates about the bore 103 when actuated by the drive system 201.

[0034] The drive system 201 rotationally actuates the gantry 210. In some embodiments, the drive system 201 includes a linear motor that can be fixed to the drive support 200 and interacts with a magnetic track (not shown) mounted on the gantry 210. In other embodiments, the drive system 201 includes another suitable drive mechanism for precisely rotating the gantry 210 about the tunnel 201. The LINAC 204 generates an MV treatment beam 230 of high energy X-rays (or electrons in some embodiments), and the EPID 205 is configured to acquire X-ray images using the treatment beam 230. The imaging X-ray source 206 is configured to direct a cone beam of X-rays (referred to herein as imaging X-rays 231) through the isocenter 203 of the RT system 100 to the X-ray imager 207, and the isocenter 203 generally corresponds to the location of the target area 209 to be treated. In Figure 2 In the illustrated embodiment, the X-ray imager 207 is depicted as a planar device, while in other embodiments, the X-ray imager 207 may have a curved configuration.

[0035] The X-ray imager 207 receives the imaging X-rays 231 and generates a suitable projection image therefrom. According to certain embodiments, such a projection image may then be used to construct or update partial imaging data of the imaging data of the digital volume corresponding to the three-dimensional (3D) region including the target region 209. That is, a 3D image of such a 3D region is reconstructed according to the projection image. In an embodiment, cone beam computed tomography (CBCT) and / or digital tomosynthesis (DTS) may be used to process the projection images generated by the X-ray imager 207. CBCT is typically used to acquire projection images over a relatively long acquisition arc (e.g., over a rotation of 180° or more of the gantry 210). As a result, a high-quality 3D reconstruction of the imaging volume may be generated. CBCT is typically used at the beginning of a radiotherapy session to generate a 3D reconstruction of the setting. For example, CBCT may be used immediately before applying the treatment beam 230 to generate a 3D reconstruction to confirm that the target region 209 has not moved or has not changed shape.

[0036] In certain embodiments, partial data reconstruction may be performed by the RT system 100 during a portion of the IGRT procedure where partial image data is used to generate a 3D reconstruction of the target volume 209. For example, while the treatment beam 230 is directed to the isocenter 203 while the gantry 210 rotates through the treatment arc, DTS image acquisition may be performed to generate image data of the target volume 209. Because DTS image acquisition is performed over a relatively short acquisition arc (e.g., between about 10° and 60°), near real-time feedback on the shape and position of the target volume 209 may be provided by DTS imaging during the IGRT procedure. Alternatively, CBCT may be employed during a portion of the IGRT procedure to generate a 3D reconstruction of the target volume 209. According to various embodiments described below, for either case, a higher frame rate X-ray image may be generated with little or no increase in image lag. Such a higher frame rate X-ray image is highly beneficial for generating accurate image data for the target volume 209, whether for CBCT or DTS image acquisition.

[0037] In some embodiments, the X-ray imager 207 includes a glass plate having a matrix or array of pixel detector elements or pixels formed thereon, each of which converts an incident X-ray photon into an electrical charge. In embodiments where the X-ray imager 207 is configured as an indirect flat panel detector, a scintillator material in the X-ray imager 207 is excited by incident X-rays and emits light, which is detected by a plurality of photodiodes. Each photodiode generates a signal (e.g., an accumulated voltage proportional to the intensity of the incident light) for a different pixel that will ultimately become a digital image. An encoder included in the X-ray imager 207 then interprets each of these voltages and assigns a value proportional to that voltage to each voltage. Figure 3One such embodiment of an X-ray imager 207 is shown in FIG.

[0038] Figure 3 Schematically illustrates a cross-sectional view of an X-ray imager 207 according to an embodiment of the present disclosure. As shown, the X-ray imager 207 includes a photosensitive element and a detector circuit layer 301 formed on a substrate 302. In addition, the X-ray imager 207 includes a scintillator material layer 303 formed on the photosensitive element and the detector circuit layer 301. Also shown is an incident X-ray 309, which is incident upon being emitted by an imaging X-ray source 206 (e.g., Figure 2 The photosensitive element and detector circuit layer 301, substrate 302 and scintillator material 303 together form an X-ray imaging array 305. Note that the photosensitive element and detector circuit layer 301 is generally formed of multiple processing layers, and the X-ray imaging array 305 may include Figure 3 Additional material layers not shown.

[0039] The photosensitive element and detector circuit layer 301 generally includes a plurality of pixel detector elements 310. Each pixel detector element 310 includes a photosensitive element (such as a photodiode, a photogate, or a phototransistor) and any other circuitry suitable for operating as a pixel detector element in the X-ray imager 207. In some embodiments, the photosensitive element of the pixel detector element 310 is an amorphous silicon-based semiconductor device. The photosensitive element and detector circuit layer 301 may also include a thin film transistor (TFT) for reading out a digital signal from the pixel detector element 310. The scintillator material 303 may include one or more material layers, including but not limited to gadolinium oxysulfide (Gd2O2S:Tb), cadmium tungstate (CdWO4), bismuth germanate (Bi4Ge3O12 or BGO), cesium iodide (Csl) or cesium thallium iodide (Csl:TI), etc.

[0040] exist Figure 3 In the illustrated embodiment, the X-ray imager 207 is depicted as an indirect flat panel detector, wherein X-ray photons are converted into other photons, which in turn are detected and converted into electric charge. In other embodiments, the X-ray imager 207 may be a direct flat panel detector (FPD). In a direct FPD, incident X-ray photons are directly converted into electric charge in the amorphous selenium layer, and the resulting charge pattern is read out by suitable hardware (such as a thin film transistor (TFT) array, an active matrix array, micro plasma line addressing, etc.).

[0041] exist Figure 2In the illustrated embodiment, the RT system 100 includes a single X-ray imager and a single corresponding imaging X-ray source. In other embodiments, the RT system 100 may include two or more X-ray imagers, each having a corresponding imaging X-ray source. Thus, in such an embodiment, the RT system 100 includes a first imaging X-ray source and a corresponding X-ray imager mounted on the gantry 210, and a second imaging X-ray source and a corresponding X-ray imager mounted on the gantry 210. In such an embodiment, including multiple X-ray imagers in the RT system 100 facilitates the generation of projection images (for reconstruction of the target area) over a shorter image acquisition arc. For example, when the RT system 100 includes two X-ray imagers and corresponding X-ray sources, the image acquisition arc for acquiring projection images of a specific image quality may be approximately half of the image acquisition arc for acquiring projection images of similar image quality with a single X-ray imager and X-ray source. In addition, in such an embodiment, including multiple X-ray imagers in the RT system 100 facilitates the use of multiple X-ray source energies, because the first imaging X-ray source and the second imaging X-ray source can each operate at different energies. Alternatively, in embodiments where the RT system 100 includes a single imaging X-ray source, the single imaging X-ray source may be configured as a multi-energy source.

[0042] The projection images generated by the X-ray imager 207 are used to construct imaging data for a digital volume of the patient's anatomy within the 3D region including the target area. Alternatively or additionally, such projection images can be used to update portions of existing imaging data of the digital volume corresponding to the 3D region. Figure 4 One embodiment of such a digital volume is described.

[0043] Figure 4 A digital volume 400 constructed based on projection images generated by one or more X-ray imagers included in the RT system 100 according to various embodiments of the present disclosure is schematically illustrated. For example, in some embodiments, the projection images may be generated by a single X-ray imager (such as the X-ray imager 207), while in other embodiments, the projection images may be generated by multiple X-ray imagers.

[0044] The digital volume 400 includes a plurality of voxels 401 (dashed lines) of anatomical image data, wherein each voxel 401 corresponds to a different location within the digital volume 400. For clarity, Figure 4 Only a single voxel 401 is shown in FIG. The digital volume 400 corresponds to a 3D region including a target area 410. Figure 4 In FIG. 4 , the digital volume 400 is depicted as an 8×8×8 voxel cube, but in reality, the digital volume 400 typically includes many more voxels, such as Figure 4Several orders of magnitude more than shown in .

[0045] For discussion purposes, the target volume 410 may refer to a gross tumor volume (GTV), a clinical target volume (CTV), or a planning target volume (PTV) for a particular treatment. The GTV depicts the location and extent of the gross tumor, such as what can be seen or imaged; the CTV includes the GTV and an additional margin for subclinical disease spread, which is typically not imageable; and the PTV is a geometric concept designed to ensure that the appropriate radiation therapy dose is actually delivered to the CTV without adversely affecting nearby organs at risk. Therefore, the PTV is typically larger than the CTV, but in some cases may also be reduced in certain portions to provide a safety margin around organs at risk. Typically, the PTV is determined based on imaging performed prior to treatment time, and alignment of the PTV with the current position of the patient's anatomy at treatment time is facilitated by embodiments of the present disclosure.

[0046] According to various embodiments described below, image information associated with each voxel 401 of the digital volume 400 is constructed based on projection images generated by a single or multiple X-ray imagers (e.g., via a CBCT or DTS process). In some embodiments, image information associated with some or all voxels 401 of the digital volume 400 is updated by projection images generated by a single or multiple X-ray imagers via a DTS process. For example, such a DTS process can be employed after a portion of a planned treatment has begun and before the planned treatment has been completed. In this way, the position and shape of the target area 410 can be confirmed while the treatment is ongoing. Therefore, if a sufficient portion of the target area 410 is detected to extend outside the threshold area, the treatment can be suspended or modified. In this case, the modification of the treatment can be achieved by adjusting the patient position and / or the treatment beam.

[0047] Figure 5 FIG. 5 is a partial circuit diagram 500 of a photosensitive element and detector circuit layer 301 according to one embodiment of the present disclosure. The photosensitive element and detector circuit layer 301 may be included in a suitable X-ray detector panel (such as Figure 2 207). As shown, the photosensitive element and detector circuit layer 301 includes a plurality of pixel detector elements 310, each of which is communicatively coupled to a readout stage 520 via a common data line 510. Typically, the photosensitive element and detector circuit layer 301 includes an M×N matrix of pixel detector elements 310, where M is equal to the number of rows of pixel detector elements 310 and N is equal to the number of columns of pixel detector elements 310. Typically, in an X-ray detector panel, M and N have values ​​on the order of about 1000 to 4000. For clarity, in Figure 5Only a single array of pixel detector elements 310 forming a column of M pixel detector elements 310 is shown. In practice, the photosensitive element and detector circuit layer 301 comprises a total of N such arrays of pixel detector elements 310, each of which forms a row of the M rows of pixels of the X-ray detector panel.

[0048] exist Figure 5 In the illustrated embodiment, each pixel detector element 310 includes a photodiode 501 and a readout switch 502 that communicatively couples the photodiode 501 to a readout stage 520. As shown, each photodiode 501 is communicatively coupled to a bias voltage V bias , and are also communicatively coupled to data line 510 via a corresponding readout switch 502. Each readout switch 502 is typically formed as part of an associated pixel detector element 310. Alternatively, in some embodiments, each readout switch 502 is formed adjacent to an associated pixel detector element 310. In either case, readout switches 502 are typically formed as part of photosensitive element and detector circuit layer 301. For example, in some embodiments, readout switches 502 are implemented as thin film transistors (TFTs) formed on the same substrate as pixel detector elements 310.

[0049] The readout stage 520 is a readout device that is configured to read out the accumulated charge from the pixel detector elements 310 in a particular column 550 of pixel detector elements 310. The readout stage 520 reads out the particular pixel detector element 310 (e.g., the pixel detector element 310 associated with pixel M-2) when the readout switch 502 of the particular pixel detector element 310 (e.g., the readout switch 502 associated with pixel M-2) is closed and communicatively couples the pixel detector element 310 to the data line 510. In operation, for each other column (not shown) of the photosensitive element and detector circuit layer 301, a single pixel detector element 310 can be read out simultaneously by the readout stage 520 associated with that column. Thus, an entire row of pixel detector elements 310 of the photosensitive element and detector circuit layer 301 can be read out simultaneously by closing one of the readout switches 502 in each column of pixel detector elements 310 in a timely manner.

[0050] The readout stage 520 is configured to convert an analog signal (such as the charge accumulated in the pixel detector element 310) into a digital X-ray image signal. In some embodiments, the readout stage 520 includes a conversion circuit 521 for converting such a signal into a digital X-ray image signal. The conversion circuit 521 may include any technically feasible circuit suitable for performing such conversion. For example, in some embodiments, the conversion circuit 521 includes an analog-to-digital converter, an analog front end, etc. In addition, in Figure 5In the illustrated embodiment, the readout stage 520 includes a reset switch 522 configured to communicatively couple the data line 510 to a reference voltage V ref , such as ground or any other suitable reference voltage. Therefore, when the readout switch 502 of a particular pixel detector element 310 is closed and the reset switch 522 is closed, the charge currently accumulated in the particular pixel detector element 310 is discharged to the reference voltage V ref In contrast, when the readout switch 502 of a particular pixel detector element 310 is closed while the reset switch 522 is open, the charge currently accumulated in that particular pixel detector element 310 is discharged to the readout stage 520 and read by the conversion circuit 521 .

[0051] exist Figure 5 In the illustrated embodiment, the reset switch 522 is included in the readout stage 520. In other embodiments, the reset switch is still communicatively coupled to the data line 510, but is implemented external to the readout stage 520. One such embodiment is Figure 6 Shown in. Figure 6 FIG. 6 is a partial circuit diagram 600 of a photosensitive element and a detector circuit layer 301 according to another embodiment of the present disclosure. As shown in the figure, Figure 6 In the illustrated embodiment, the reset switch 622 is configured to selectively couple the data line 510 to a reference voltage V ref , but is located at some other location along the data line 510 other than the readout stage 520. In other embodiments, the reset switch is communicatively coupled to the data line 510, but is implemented via existing circuitry included in the conversion circuit 521. One such embodiment is Figure 7 Shown in. Figure 7 FIG. 7 is a partial circuit diagram 700 of a photosensitive element and a detector circuit layer 301 according to another embodiment of the present disclosure. As shown in the figure, Figure 7 In the illustrated embodiment, the reset switch 722 is configured to selectively couple the data line 510 to a reference voltage V ref , and is included in the conversion circuit 521. Alternatively, the function of the reset switch 722 is implemented by one or more components of the conversion circuit 521, and the reset switch 722 is not used to couple the data line 510 to the reference voltage V ref A dedicated switch or transistor.

[0052] Reducing image lag via pixel charge reset

[0053] According to various embodiments described herein, the acquisition of an X-ray image using an X-ray imager (such as X-ray imager 207) is performed in three phases: an illumination phase, a readout phase, and a reset phase. In the illumination phase, each pixel detector element 310 integrates the charge generated by illuminating the panel with imaging X-rays via the pixel capacitance. In the readout phase, the accumulated charge of each pixel detector element 310 is transferred to the readout stage 520 and processed. In the reset phase, residual charge is transferred from each pixel detector element 310, thereby minimizing or otherwise reducing image lag present in the next X-ray image to be acquired. The following is combined with Figure 8 One such embodiment is described.

[0054] Figure 8 8 is a timing diagram 800 schematically illustrating charge accumulation and loss in a pixel detector element 310 during an illumination phase 810, a readout phase 820, and a reset phase 830 of a single X-ray image acquisition according to an embodiment of the present disclosure. More specifically, charge accumulation and loss are illustrated for M rows of pixel detector elements 310 of an X-ray detector panel that includes an M×N matrix of pixel detector elements 310. Figure 5 describes one such row.

[0055] As shown, throughout the illumination phase 810, each of the M rows is illuminated simultaneously, and charge is accumulated in the pixel detector elements 310 of each row. In the readout phase 820, the charge accumulated in the pixel detector elements 310 of each row is read out sequentially. That is, the pixel detector elements 310 of row 1 are read out by the readout stage 520, then the pixel detector elements 310 of row 2 are read out by the readout stage 520, and so on, until all M rows of pixel detector elements 310 are read out and an X-ray image can be generated. In the reset phase 830, the residual charge remaining in the pixel detector elements 310 of each row is transferred concurrently from all rows of the X-ray detector panel. The following is combined with Fig. 9 The readout phase 820 is described in more detail below in conjunction with Fig.10 The reset phase 830 is described in more detail.

[0056] Fig. 9 8 is a schematic timing diagram 900 illustrating charge loss of two adjacent rows of pixel detector elements 310 during the readout phase 820 according to an embodiment of the present disclosure. Fig. 9 , a portion of a readout phase 820 is shown for a first array of pixel detector elements 310 (row M-4) and an adjacent second array of pixel detector elements 310 (row M-3). During the readout phase 820, at a first readout time interval t read1The accumulated charge in each of the pixel detector elements 310 in row M-4 is read during the first readout time interval t read1 The second readout time interval after t read2 During this period, the accumulated charge in each of the pixel detector elements 310 in row M-3 is read.

[0057] First readout time interval t read1 The first readout time interval t begins after the accumulated charge in each of the pixel detector elements 310 in the previous row (e.g., row M-5, not shown) has been read out, because the photosensitive element and detector circuit layer 301 is typically connected to a readout stage 520 configured to read out one row of pixel detector elements 310 at a time. Similarly, the second readout time interval t read2 The accumulated charge in each of the pixel detector elements 310 in row M-4 has been read out and the first readout time interval t read1 After it's over.

[0058] In the first readout time interval t read1 During this period, the readout switch 502 for each pixel detector element 310 in row M-4 is closed (at Fig. 9 ) and accumulates charge Q M-4 In the first readout time interval t read1 The magnitude of the charge is increased from the initial charge value Q 0 M-4 Reduced to the residual charge value Q Rem M-4 Then, the readout switch 502 for each pixel detector element 310 in row M-4 is opened (at Fig. 9 OFF state in the row M-4), and no more charge is read out from the pixel detector element 310 in row M-4. Typically, the remaining charge value Q Rem M-4 The value is the first read time interval t read1 duration, the time-dependent release of trapped charge and the pixel time constant τ pix function, constant τ pix The same is true for each photodiode 501 of the X-ray imager. For example, in some embodiments, the residual charge value Q can be determined based on Equation 1: Rem M-4 Values:

[0059] Q Rem M-4 =Q 0 M-4 ×exp(-tread1 / τ pix )+Q trap M-4 (t) (1)

[0060] Similarly, at the second readout time interval t read2 During this period, the readout switch 502 for each pixel detector element 310 in row M-3 is closed (at Fig. 9 ) and accumulates charge Q M-3 At the second readout time interval t read2 The value of the internal charge is from the initial charge value Q 0 M-3 Reduced to the residual charge value Q Rem M-3 Then, the readout switch 502 for each pixel detector element 310 in row M-3 is opened (at Fig. 9 OFF state in row M-3), and no more charge is read out from the pixel detector elements 310 in row M-3. This process continues sequentially through the remaining rows of the X-ray imager.

[0061] Once the associated readout switch 502 is opened, the charge will not escape from the photodiode 501. Therefore, in a conventional X-ray imager, the residual charge value Q of a row is Rem The value of Q corresponds to an image lag for each of the pixel detector elements 310 in the row, since this residual charge is present at the start of the subsequent illumination phase 810. As a result, the image quality suffers. Alternatively, the duration of each readout time interval may be increased so that the residual charge value Q Rem The magnitude of is irrelevant. In the latter case, the panel readout time of the X-ray imager is greatly slowed down, because the panel readout time increases based on the following relationship: (increase in readout time interval) × (number of rows of pixel detector elements). In addition, because the residual charge value Q Rem decays at an exponential rate during readout, so a relatively large increase in the readout time interval is required to produce a residual charge value Q Rem In contrast, according to various embodiments described herein, the reset phase 830 causes the residual charge value Q of each pixel of the X-ray detector panel to be Rem can be significantly reduced in the relatively short time interval before the next illumination phase 810. Thus, image lag can be prevented without slowing down the panel readout time beyond the duration of the reset phase 830.

[0062] Fig.10 1 is a schematic timing diagram 1000 illustrating charge loss for two adjacent rows of pixel detector elements 310 during the reset phase 830 according to an embodiment of the present disclosure. Fig.10 , a portion of a reset phase 830 is shown for a first array (row M-4) of pixel detector elements 310 and a second array (row M-3) of adjacent pixel detector elements 310. The reset phase 830 begins after the readout phase 820 for all M rows of the X-ray detector panel has been completed.

[0063] During the reset phase 830, the residual charge in each of the pixel detector elements 310 (eg, the residual charge in a single reset time interval t reset During the reset time interval t reset During this period, the residual charge in each of the pixel detector elements 310 in row M-4 is transferred from the pixel detector element, for example, to ground or some other suitable reference voltage V ref In addition, during the reset time interval t reset During this period, the residual charge in each of the pixel detector elements 310 in row M-3 is transferred from the pixel detector element to the reference voltage V ref In addition, during the reset time interval t reset During this time, the residual charge in each of the pixel detector elements 310 in the remaining rows (not shown) of the X-ray detector panel is transferred from the pixel detector element to the reference voltage V ref Therefore, during the reset time interval t reset During this time, residual charge from some or all of the M rows is concurrently transferred from pixel detector element 310 to reference voltage V ref As a result, the readout time interval t for each row of pixel detector elements 310 is increased by read The magnitude of the image lag associated with each pixel detector element 310 may be reduced more quickly than the duration of the image lag.

[0064] like Fig.10 As shown, during the reset time interval t reset Initially, the residual charge in each particular pixel detector element 310 in row M-4 is equal to the residual charge value Q Rem M -4, and at the reset time interval t reset At the end, the residual charge in each particular pixel detector element 310 in row M-4 is equal to the final charge value Q of the particular pixel detector element 310. Fin M-4 Similarly, during the reset time interval t resetInitially, the residual charge in each particular pixel detector element 310 in row M-3 is equal to the residual charge value Q of that particular pixel detector element 310. Rem M-3 , and at the reset time interval t reset At the end, the residual charge in each particular pixel detector element 310 in row M-3 is equal to the final charge value Q of the particular pixel detector element 310. Fin M-3 . Similar to the residual charge value Q Rem , in some embodiments, the final charge value Q of each pixel detector element 310 in a particular row (e.g., row M-3) Fin It can be determined based on Equation 2:

[0065] Q Fin M-3 =Q 0 M-3 ×exp(-t reset / τ pix )+Q trap M-4 (t)(2)

[0066] Therefore, the reset time interval t in the reset phase 830 is reset During this period, the Q of the pixel detector element 310 Fin The value of Q Rem The value of read1 ) decays at substantially the same rate during the reset phase. However, during the reset phase 830, charge is transferred simultaneously from some, most, or all of the M rows of pixel detector elements 310 of the X-ray detector panel. Therefore, the reset time interval t is increased. reset The duration of can significantly reduce the magnitude of the accumulated charge in each of the pixel detector elements 310 of the X-ray detector panel (from Q Rem To Q Fin ) without significantly increasing the panel readout time of the X-ray detector panel. In contrast, the same reduction in accumulated charge (down to Q ) is achieved by increasing the duration of each of the readout time intervals in the readout phase 820. Fin ), the panel readout time of the X-ray detector panel is greatly increased. For example, in one embodiment, the reset time interval t of the reset phase 830 is reset is set to be three times longer than the readout time interval of the readout phase 820. Therefore, the panel readout time is increased by a factor of 3 (the readout time interval t reset). To achieve the same reduction in accumulated charge in the pixel detector elements 310 of an X-ray detector panel without using the reset phase 830, the panel readout time is four times, which significantly slows down the imaging frame rate of the X-ray detector.

[0067] Fig.11 A flow chart of a method for acquiring X-ray image data in an X-ray detector panel according to one or more embodiments of the present disclosure is set forth. The method may include one or more operations, functions, or actions as shown in one or more of blocks 1111-1135. Although the blocks are shown in a sequential order, the blocks may be executed in parallel and / or in an order different from that described herein. In addition, various blocks may be combined into fewer blocks, divided into additional blocks, and / or eliminated based on the desired implementation. Although combined Figure 1-Figure 10 The method is described with reference to a system of the present invention, but those skilled in the art will appreciate that any suitably configured radiation therapy system is within the scope of the present disclosure. The control algorithms for the method steps may be implemented in whole or in part as software or firmware implemented logic, and / or hardware implemented logic circuits. In addition, the control algorithms for the method steps may be implemented in whole or in part by the treatment control computer 106 (e.g., the RT system 100) of the present invention. Figure 1 ), including a controller in the X-ray imager 207 (such as Figure 2 as shown), any other suitable controller associated with the RT system 100, or any combination thereof.

[0068] The method 1100 begins at step 1111, where the treatment control computer 106, a controller included in the X-ray imager 207, or any other suitable controller associated with the RT system 100 prepares the X-ray imager 207 for the irradiation phase 810 and the acquisition of an X-ray image. In some embodiments, the controller causes the charge currently accumulated in the pixel detector elements 310 of the X-ray imager 207 to be transferred to a ground or other reference voltage V ref For example, the controller closes the readout switch 502 and the reset switch 522 of each pixel detector element 310 for a certain time interval to reduce the accumulated charge in the pixel detector element 310. The controller then opens the readout switch 502 and the reset switch 522 of each pixel detector element 310.

[0069] In step 1112 , the controller begins the irradiation phase 810 . For example, in some embodiments, the controller causes the imaging X-ray source 206 to direct the imaging X-rays 231 through the isocenter 203 of the RT system 100 to the X-ray imager 207 .

[0070] In step 1113 , the X-ray imager 207 receives the imaging X-rays 231 , and charge is accumulated in some or all of the pixel detector elements 310 .

[0071] In step 1114, the controller ends the irradiation phase 810. For example, in some embodiments, the controller causes the imaging X-ray source 206 to stop directing the imaging X-rays 231 to the X-ray imager 207.

[0072] In step 1121, the controller selects a pixel array for readout from the M pixel arrays of the X-ray imager 207. In some embodiments, each of the M pixel arrays is configured as a row of pixel detector elements 310. In other embodiments, each of the M pixel arrays is configured as a column of pixel detector elements 310. In other embodiments, each of the M pixel arrays is configured as any other group of pixel detector elements 310 that are read out simultaneously by the readout stage 520, such as a group of pixel detector elements 310 located in a particular area of ​​the X-ray imager 207. For clarity, the method 1100 is described herein in terms of rows of pixel detector elements 310 included in the X-ray imager 207, but the method 1100 is equally applicable to any other suitable pixel array configuration of pixel detector elements 310, such as columns or other groups of pixel detector elements 310.

[0073] In step 1122, the controller causes the accumulated charge from the pixel detector element 310 in the selected pixel array to be read out during the readout interval. For example, in some embodiments, the controller causes the readout switch 502 of each pixel detector element 310 in the selected pixel array to be switched on during the readout interval t read The internal circuit is closed and the reference voltage V ref The reset switch 522 of step 1122 remains open. As a result, the charge currently accumulated in each pixel detector element 310 in the selected pixel array is discharged to the readout stage 520, read by the conversion circuit 521, and processed as part of the current X-ray image being acquired. During the implementation of step 1122, the accumulated charge present in each pixel detector element 310 in the selected pixel array is increased from the initial charge value Q at the rate described by the previously given equation 1. 0 Reduced to the residual charge value Q Rem Note that for each pixel detector element 310, the initial charge value Q 0 are usually different. Similarly, the residual charge value Q Rem is the initial charge value Q 0 is a function of , and is therefore typically different for each pixel detector element 310 .

[0074] In step 1123 , the controller determines whether there are remaining pixel arrays to be read out. If yes, the method 1100 returns to step 1121 ; if no, the method 1100 proceeds to step 1131 .

[0075] In step 1131, the controller begins the reset phase 830. For example, in some embodiments, the controller causes the reset switch 522 to close and the data line 510 to be communicatively coupled to V ref .

[0076] In step 1132, the controller selects one or more pixel arrays in the X-ray detector 207 to be reset. Fig.10 In the embodiment of the reset phase 830 shown, the controller selects a plurality of pixel arrays of the X-ray detector 207, namely all M rows of pixel detector elements 310. In such an embodiment, the residual charge values ​​Q of all M rows of pixel detector elements 310 are Rem At the same time. Alternatively, in some embodiments, a staged reset of the pixel array is performed during the reset phase. In such an embodiment, the peak current exposed to the readout stage 522 during the reset phase 830 can be reduced by preventing all pixel arrays of the X-ray detector 207 from being coupled to the data line 510 at the same time. Therefore, in some embodiments, in step 1132, the controller selects a portion of the total M rows of the pixel array (e.g., one-tenth of the M rows) to start resetting in step 1133. In such an embodiment, a subsequent portion is selected to begin being reset while the previously selected portion continues to be reset. One such embodiment is Fig. 12A and Fig. 12B Shown in.

[0077] Fig. 12A is a timing diagram 1200 schematically illustrating charge loss for different rows of pixel detector elements 310 during a reset phase 1230 according to an embodiment of the present disclosure. Fig. 12B A portion of a reset phase 1230 of a first array of pixel detector elements 310 (row M- 4 ) and an adjacent second array of pixel detector elements 310 (row M- 3 ) is shown in accordance with an embodiment of the present disclosure.

[0078] The reset phase 1230 is similar to Figure 8The reset phase 1230 of the X-ray detector panel is shown in FIG. 1230 because the residual charge in each pixel detector element 310 in the plurality of rows of the X-ray detector panel is transferred from the pixel detector elements concurrently during the reset phase 1230. That is, during at least a portion of the reset phase 1230, the plurality of rows of the X-ray detector panel are simultaneously undergoing reset, and the accumulated charge is being transferred from all pixel detector elements 310 associated with the plurality of rows simultaneously. However, during the reset phase 1230, the reset time interval t reset The reset of the M rows of pixel detector elements 310 is therefore performed in stages throughout the reset phase 1230. Therefore, when a row of pixel detector elements 310 is initially coupled to the data line 510, the large initial current generated in the data line 510 does not appear simultaneously in all rows of pixel detector elements 310. Instead, multiple rows of pixel detector elements 310 are initially coupled to the data line 510 and the readout stage 522 at different times during the reset phase 1230.

[0079] exist Fig. 12A and Fig. 12B In the illustrated embodiment, each of the M rows of pixel detector elements 310 is initially coupled to the readout stage 522 at a different time during the reset phase 1230. Thus, in such an embodiment, each row of pixel detector elements 310 is associated with a different reset initiation time. Fig. 12B As shown, during the reset start time t reset,M-4 , starts resetting the pixel detector element 310 of row M-4, and at the reset start time t reset,M-3 , initiates the reset of the pixel detector element 310 of row M-3. In some embodiments, t reset,M-3 In time with the reset start time t reset,M-4 Separate the segments by time interval t stage In some embodiments, the segment time interval t stage The time intervals t are generally uniform between successive rows of pixel detector elements 310, while in other embodiments, the time intervals t stage can be varied between different consecutive rows of pixel detector elements 310. Typically, the segment time interval t stage The duration of the reset phase 1230 is relatively small compared to the duration of the reset phase 1230. Therefore, in embodiments where the reset of the M rows of pixel detector elements 310 is staged throughout the reset phase 1230, the total duration of the reset phase 1230 is not significantly increased.

[0080] exist Fig. 12A and Fig. 12BIn the illustrated embodiment, each of the M rows of pixel detector elements 310 is initially coupled to the readout stage 522 at a unique time in the reset phase 1230. Alternatively, in some embodiments, a first group of multiple rows of pixel detector elements 310 are initially coupled to the readout stage 522 at the same time in the reset phase 1230, a second group of multiple rows of pixel detector elements 310 are initially coupled to the readout stage 522 at a later time in the reset phase 1230, a third group of multiple rows of pixel detector elements 310 are initially coupled to the readout stage 522 at another later time in the reset phase 1230, and so on. For example, in one such embodiment, each such group of multiple rows of pixel detector elements 310 includes 10% of the M rows of pixel detector elements 310, and each such group includes pixel detector elements 310 of a different row than the other rows of pixel detector elements 310. Thus, in such an embodiment, the reset phase 1230 includes 10 different reset initiation times, rather than one reset initiation time for each row of pixel detector elements.

[0081] return Fig.11 In step 1133, the controller causes one or more rows of pixel detector elements 310 selected in step 1132 to be reset. For example, in some embodiments, the controller causes reset switch 522 to close, thereby communicatively coupling data line 510 to reference voltage V ref In addition, the controller closes the readout switch 502 of each pixel detector element 310 in the selected row or rows, thereby communicatively coupling such pixel detector element 310 to the data line 510. As a result, the charge present in the pixel detector element 310 (e.g., the residual charge value Q Rem ) begins to be transferred to the reference voltage V ref . Thus, the charge present in pixel detector element 310 decreases at the rate indicated by Equation 2 given previously.

[0082] In an embodiment where all M rows of pixel detector elements 310 are selected in step 1131, all pixel detector elements 310 of the X-ray imager 207 are reset simultaneously. In such an embodiment, during the reset time interval t reset After passing, method 1100 proceeds to step 1133.

[0083] In an embodiment where a segmented reset of the pixel array is performed during the reset phase, a single row of pixel detector elements 310 is reset beginning in step 1132, or a group of multiple rows of pixel detector elements 310 are reset beginning in step 1132. In such an embodiment, at segmented time intervals t stage After has passed, method 1100 proceeds to step 1133.

[0084] In step 1134, the controller determines whether there are remaining pixel arrays that have not yet begun to be reset. If so, the method returns to step 1132; if not, the method 1100 proceeds to step 1135. Note that in an embodiment in which a segmented reset of the pixel arrays is performed during the reset phase, the method 1100 may return to step 1132 while some pixel arrays of the X-ray imager 207 are currently being reset. That is, the controller may reset the pixel arrays during the reset time interval t for some or all pixel arrays that have been selected for reset. reset Before it has passed, execute step 1134.

[0085] In step 1135, the controller determines whether the reset has been completed for all pixel arrays. That is, the controller determines whether the reset time interval t reset If not, the method 1100 returns to step 1135; if yes, the method 1100 proceeds to step 1136.

[0086] In step 1136, the controller ends the reset phase. For example, in some embodiments, the controller opens the reset switch 522 and opens the readout switch 502 of each pixel detector element 310 of the X-ray imager 207. When further acquisition of X-ray images is planned, the method 1100 returns to step 1112.

[0087] Implementation of method 1100 enables the reset time interval t reset The remaining charge in the pixel detector element 310 is significantly reduced (e.g., from Q Rem To Q Fin As mentioned above, the reset time interval t is shorter than the duration of the read time interval. reset can be chosen to be relatively short. In conventional X-ray imagers, an equivalent charge reduction can only be achieved over a much larger time interval, i.e., (reset time interval t reset )×(number of rows of pixel detector elements 310). Since the number of rows of pixel detector elements 310 can be on the order of 1000 to 4000, the panel readout time of a conventional X-ray imager can be significantly increased, thereby preventing rapid image acquisition.

[0088] Synchronize resetting with application of treatment beam

[0089] In practice, the treatment beam in an RT system typically generates a large amount of scattered radiation in all directions, including radiation emanating from the patient, the treatment table, and machine components. As a result, a large amount of MV scatter may be incident on the X-ray imager (e.g., Figure 2In some cases, the amount of such X-ray scatter can even exceed the amount of imaging X-rays. Accordingly, in some embodiments, a reset phase (such as reset phase 830 or reset phase 1230) is timed to coincide with the application of a treatment beam (such as treatment beam 230). As described herein, the X-ray imager 207 is insensitive to radiation during the reset phase. That is, the pixel detector elements 310 of the X-ray imager 207 do not accumulate charge during such a reset phase, even when the treatment beam 230 produces significant X-ray scatter incident on the X-ray imager 207. As a result, implementing the reset phase to coincide with a burst or other application of the treatment beam 231 reduces the noise typically caused by the treatment beam 230. Fig.13 One such embodiment is shown in .

[0090] Fig.13 1 is a timing diagram 1300 schematically illustrating the relative timing of the irradiation phase 1310, the readout phase 1320, and the reset phase 1330 with respect to the delivery of the treatment beam pulse 1350 according to an embodiment of the present disclosure. As shown, the KV (imaging) beam pulse 1360 occurs during each irradiation phase 1310, and the delivery of the treatment beam pulse 1350 is timed to occur during each reset phase 1330. As described above, the pixel detector elements 310 of the X-ray imager 207 do not accumulate charge during the reset phase. Therefore, even though scattered X-rays from the treatment beam pulse 1350 may strike the scintillator layer of the X-ray imager 207, noise is not added to the image data generated by the X-ray imager 207.

[0091] Although embodiments are described herein with respect to an X-ray imager included in an RT system, the embodiments are equally applicable to other X-ray imaging systems. For example, the embodiments may also be implemented in a handheld or portable flat panel X-ray detector (FPD), a statically mounted FPD, an X-ray imager configured for dynamic X-ray imaging (such as fluoroscopic imaging), etc. In addition, the embodiments may be used to detect intra-subject motion prior to treatment or pre-treatment imaging.

[0092] The description of various embodiments is for illustrative purposes only, but is not intended to be exhaustive or to limit the disclosed embodiments. Many modifications and variations will be apparent to those of ordinary skill in the art without departing from the scope and spirit of the described embodiments.

[0093] Aspects of the present embodiment may be embodied as a system, method, or computer program product. Accordingly, aspects of the present disclosure may take the form of a complete hardware embodiment, a complete software embodiment (including firmware, resident software, microcode, etc.), or an embodiment combining software aspects and hardware aspects (which may be collectively referred to herein as "circuits," "modules," or "systems"). In addition, aspects of the present disclosure may take the form of a computer program product embodied in one or more computer-readable media having computer-readable program code thereon.

[0094] Any combination of one or more computer-readable media can be utilized. Computer-readable media can be computer-readable signal media or computer-readable storage media. Computer-readable storage media can be, for example, but not limited to, electronic, magnetic, optical, electromagnetic, infrared or semiconductor systems, devices or equipment, or any suitable combination of the foregoing. More specific examples (non-exhaustive list) of computer-readable storage media will include the following: an electrical connection with one or more wires, a portable computer disk, a hard disk, a random access memory (RAM), a read-only memory (ROM), an erasable programmable read-only memory (EPROM or flash memory), an optical fiber, a portable compact disk read-only memory (CD-ROM), an optical storage device, a magnetic storage device, or any suitable combination of the foregoing. In the context of this article, a computer-readable storage medium can be any tangible medium that can contain or store a program used by or in conjunction with an instruction execution system, device or device.

[0095] While various aspects and embodiments are disclosed herein, other aspects and embodiments will be apparent to those skilled in the art. The various aspects and embodiments disclosed herein are for purposes of illustration and are not intended to be limiting, with the true scope and spirit being indicated by the following claims.

Claims

1. A method for acquiring an X-ray image of a target area, the method comprising: include: transferring accumulated charge from each pixel in a first pixel array of the X-ray detector panel to a readout device during a first readout interval, wherein after said transferring, a residual charge remains in each pixel in said first pixel array; transferring the accumulated charge from each pixel in a second pixel array of the X-ray detector panel to the readout device during a second readout interval after the first readout interval, wherein a residual charge remains in each pixel in the second pixel array after the accumulated charge is transferred from each pixel in the second pixel array; and concurrently transferring at least a portion of the residual charge from each pixel in the first pixel array and at least a portion of the residual charge from each pixel in the second pixel array during a reset interval following the second readout interval, Wherein a start time of the reset interval is selected to occur before application of the treatment beam, and an end time of the reset interval is selected to occur after application of the treatment beam.

2. The method according to claim 1, further comprising: include: During the reset interval, at least a portion of the residual charge from each pixel in each pixel array of the X-ray detector panel is concurrently transferred.

3. The method according to claim 1 or 2, in, Transferring the residual charge from each pixel in the first pixel array includes transferring the residual charge from each pixel in the first pixel array to the readout device.

4. The method according to claim 3, in, Transferring the residual charge from each pixel in the first pixel array to the readout device includes electrically coupling the first pixel array to a readout stage of the readout device.

5. The method according to any one of claims 1 to 4, in, Transferring the residual charge from each pixel in the first pixel array includes electrically coupling each pixel in the first pixel array to a reference voltage.

6. The method according to any one of claims 1 to 5, in, The first pixel array includes a first row of pixels, and the second pixel array includes a second row of pixels.

7. The method of claim 1, wherein the treatment beam is a megavoltage treatment beam directed to the target area.

8. The method according to claim 1, in, Concurrently transferring at least a portion of the residual charge from each pixel in the first pixel array and at least a portion of the residual charge from each pixel in the second pixel array comprises: communicatively coupling the first pixel array to a reference voltage so that the residual charge from each pixel in the first pixel array begins to be transferred to the reference voltage; and While the residual charge from each pixel in the first pixel array is being transferred to the reference voltage, the second pixel array is communicatively coupled to a reference voltage such that the residual charge from each pixel in the second pixel array begins to be transferred to the reference voltage.

9. The method according to claim 1, in, Concurrently transferring at least a portion of the residual charge from each pixel in the first pixel array and at least a portion of the residual charge from each pixel in the second pixel array comprises: communicatively coupling a first set of pixel arrays including the first pixel array to a reference voltage such that the residual charge from each pixel in the first set of pixel arrays begins to be transferred to the reference voltage; and While the residual charge from each pixel in the first group of pixel arrays is being transferred to the reference voltage, a second group of pixel arrays including the second pixel array is communicatively coupled to the reference voltage so that the residual charge from each pixel in the second group of pixel arrays begins to be transferred to the reference voltage.

10. The method according to any one of claims 1 to 9, further comprising: include: preparing each pixel in the first pixel array and each pixel in the second pixel array to accumulate charge in response to incident X-rays prior to the first readout interval; as well as Imaging X-rays are directed to the X-ray detector panel.

11. A device, include: an imaging X-ray source configured to direct imaging X-rays through the target region and toward the X-ray detector panel; as well as The controller is configured as: During a first readout interval, the accumulated charge from each pixel in the first pixel array of the X-ray detector panel is transferred to a readout device, wherein After the transfer, a residual charge remains in each pixel in the first pixel array; causing, after the first readout interval, during a second readout interval, accumulated charge from each pixel in a second pixel array of the X-ray detector panel to be transferred to the readout device, wherein after transferring the accumulated charge from each pixel in the second pixel array, a residual charge remains in each pixel in the second pixel array; as well as causing at least a portion of the residual charge to be transferred from each pixel in the first pixel array during a reset interval following the second readout interval, and concurrently causing at least a portion of the residual charge to be transferred from each pixel in the second pixel array, Wherein a start time of the reset interval is selected to occur before application of the treatment beam, and an end time of the reset interval is selected to occur after application of the treatment beam.

12. The device according to claim 11, further comprising: include: During the reset interval, at least a portion of the residual charge from each pixel in each pixel array of the X-ray detector panel is concurrently transferred.

13. The device according to claim 11 or 12, in, Transferring the residual charge from each pixel in the first pixel array includes transferring the residual charge from each pixel in the first pixel array to the readout device.

14. The device according to claim 13, in, Transferring the residual charge from each pixel in the first pixel array to the readout device includes electrically coupling the first pixel array to a readout stage of the readout device.

15. The device according to any one of claims 11 to 14, in, Transferring the residual charge from each pixel in the first pixel array includes electrically coupling each pixel in the first pixel array to a reference voltage.

16. The device according to any one of claims 11 to 15, in, The first pixel array includes a first row of pixels, and the second pixel array includes a second row of pixels.

17. The apparatus of claim 11, wherein at least a portion of the residual charge from each pixel in the first pixel array and at least a portion of the residual charge from each pixel in the second pixel array are concurrently transferred. include: communicatively coupling the first pixel array to a reference voltage so that residual charge from each pixel in the first pixel array begins to be transferred to the reference voltage; as well as While residual charge from each pixel in the first pixel array is being transferred to the reference voltage, the second pixel array is communicatively coupled to the reference voltage such that residual charge from each pixel in the second pixel array begins to be transferred to the reference voltage.

18. The device according to claim 11, in, Concurrently transferring at least a portion of the residual charge from each pixel in the first pixel array and at least a portion of the residual charge from each pixel in the second pixel array comprises: communicatively coupling a first set of pixel arrays including the first pixel array to a reference voltage so that residual charge from each pixel in the first set of pixel arrays begins to be transferred to the reference voltage; and While the residual charge from each pixel in the first group of pixel arrays is being transferred to the reference voltage, a second group of pixel arrays including the second pixel array is communicatively coupled to the reference voltage so that the residual charge from each pixel in the second group of pixel arrays begins to be transferred to the reference voltage.

Images