Patent Description:
So that the prescribed dose is correctly supplied to the planning target volume (i.e., the target tissue) during radiation therapy, the patient should be correctly positioned relative to the linear accelerator that provides the radiation therapy. Typically, dosimetric and geometric data are checked before and during the treatment, to ensure correct patient placement and that the administered radiotherapy treatment matches the previously planned treatment. This process is referred to as image guided radiation therapy (IGRT), and involves the use of an imaging system to view target tissues during or prior to radiation treatment delivery to the planning target volume. IGRT incorporates imaging coordinates from the treatment plan to ensure the patient is properly aligned for treatment in the radiation therapy device.

<CIT> shows an X-ray image sensor with two groups of pixels which are read out sequentially and resetting is done with multiple rows in parallel.

<CIT> shows an Xray image sensor with the capability of adding a predefined reset voltage to the pixels while resetting them.

<CIT> and X-ray image sensor where multiple or all rows are reset simultaneously.

The main embodiments of the invention are defined by the appended independent claims. The appended dependent claims define further embodiments.

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

The foregoing and other features of the present disclosure will become more fully apparent from the following description and appended claims, taken in conjunction with the accompanying drawings. These drawings depict only several embodiments in accordance with the disclosure and are, therefore, not to be considered limiting of its scope. The disclosure will be described with additional specificity and detail through use of the accompanying drawings.

In the following detailed description, reference is made to the accompanying drawings, which form a part hereof. In the drawings, similar symbols typically identify similar components, unless context dictates otherwise. It will be readily understood that the aspects of the disclosure, as generally described herein, and illustrated in the figures, can be arranged, substituted, combined, and designed in a wide variety of different configurations, all of which are explicitly contemplated and make part of this disclosure.

Image guided radiation therapy (IGRT) is used to treat tumors in areas of the body that are subject to voluntary movement, such as the lungs, or involuntary movement, such as organs affected by peristalsis. IGRT involves the use of an imaging system to view target tissues (also referred to as the "target volume") while radiation treatment is delivered thereto. In IGRT, image-based coordinates of the target volume from a previously determined treatment plan are compared to 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 motion or deformation of the target volume relative to the radiation therapy system can be detected. Consequently, dose limits to organs at risk are accurately enforced based on the daily position and shape, and the patient's position and/or the treatment beam can be adjusted to more precisely target the radiation dose to the tumor. For example, in pancreatic tumor treatments, organs at risk include the duodenum and stomach. The shape and relative position of these organs at risk with respect to the target volume can vary significantly from day-to-day. Thus, accurate adaption to the shape and relative position of such organs at risk enables dose sparing to those organs at risk and escalation of the dose to the target volume and better therapeutic results.

In some conventional IGRT radiation systems, motion of soft tissues is detected during application of the treatment beam via fiducial markers, such as gold seeds. However, the use of fiducial markers has numerous drawbacks, particularly the invasive surgical procedures required for placement of the markers. Specifically, the laproscopic insertion of fiducial markers requires additional time and clinical resources, such as an operating room, anesthesia, antibiotics, and the participation of numerous additional medical specialists.

Alternatively, in some conventional IGRT radiation systems, motion of soft tissues is detected during application of the treatment beam via magnetic resonance imaging (MRI). However, MRI-based IGRT also has drawbacks. First, MRI-based IGRT systems are generally larger, more complex, and more expensive than radiation therapy systems that employ X-ray imaging. Second, detecting motion or deformation of the target volume via MRI generally involves monitoring images associated with a 2D slice that passes through the target volume. As a result, target volume motion or deformation that occurs anywhere outside of (or perpendicular to) the 2D slice being monitored is difficult to detect, which can significantly impact the accuracy of the radiation dose being applied.

Alternatively, in some conventional IGRT radiation systems, motion of soft tissues is detected during application of treatment X-rays via imaging X-rays that are also directed through the target volume. For example, volumetric image data for the target volume can be reconstructed based on X-ray projection images of the target volume that are generated with a computed tomography (CT) or cone-beam CT (CBCT) process. In a CT or CBCT process, a plurality of X-ray projection images are generated by the imaging X-rays passing though the target volume and onto an X-ray detector panel or other X-ray imaging device. Generally, in IGRT applications, faster CT or CBCT acquisition is beneficial, since faster acquisition of target volume images enables faster detection of motion or deformation of the target volume and/or changes in the surrounding organs at risk.

The speed of CT or CBCT acquisition is strongly dependent on panel readout time of the X-ray detector panel generating the X-ray projection images of the target volume and surrounding organs at risk. Because all pixels in one row of an 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. Thus, application of a shorter pixel readout time in an X-ray detector panel can significantly reduce panel readout time. However, a shorter pixel readout time in certain X-ray detector panels, such as amorphous silicon-based panels, necessarily causes greater image lag in X-ray images generated by such X-ray detector panels. Image lag is the carryover of charge associated with the pixels of one X-ray image to the pixels of a subsequent X-ray image, and can cause significant image artifacts. Consequently, in implementing faster X-ray image acquisition, there is a well-known trade-off between imaging framerate and image lag.

In light of the above, there is a need in the art for improved systems and techniques for increasing imaging framerate in a radiation therapy system without increasing image lag in the resultant X-ray images. One such embodiment is illustrated in <FIG>.

<FIG> is a perspective view of a radiation therapy system <NUM> that can beneficially implement various aspects of the present disclosure. Radiation therapy (RT) system <NUM> is a radiation system configured to detect intra-fraction motion in near-real time using X-ray imaging techniques. Thus, RT system <NUM> is configured to provide stereotactic radiosurgery and precision radiotherapy for lesions, tumors, and conditions anywhere in the body where radiation treatment is indicated. As such, RT system <NUM> can include one or more of a linear accelerator (LINAC) that generates a megavolt (MV) treatment beam of high energy X-rays, a kilovolt (kV) X-ray source, an X-ray imager, and, in some embodiments, an MV electronic portal imaging device (EPID). By way of example, radiation therapy system <NUM> is described herein configured with a circular gantry. In other embodiments, radiation therapy system <NUM> can be configured with a C-gantry capable of infinite rotation via a slip ring connection, a ring gantry with a slip ring, a C-gantry with a wind-up configuration, and the like.

Generally, RT system <NUM> is capable of kV imaging of a target volume during application of an MV treatment beam, so that an IGRT process can be performed using X-ray imaging rather than MRI. RT system <NUM> may include one or more touchscreens <NUM>, couch motion controls <NUM>, a bore <NUM>, a base positioning assembly <NUM>, a couch <NUM> disposed on base positioning assembly <NUM>, and an image acquisition and treatment control computer <NUM>, all of which are disposed within a treatment room. RT system <NUM> further includes a remote control console <NUM>, which is disposed outside the treatment room and enables treatment delivery and patient monitoring from a remote location. Base positioning assembly <NUM> is configured to precisely position couch <NUM> with respect to bore <NUM>, and motion controls <NUM> include input devices, such as button and/or switches, that enable a user to operate base positioning assembly <NUM> to automatically and precisely position couch <NUM> to a predetermined location with respect to bore <NUM>. Motion controls <NUM> also enable a user to manually position couch <NUM> to a predetermined location. In some embodiments, RT system <NUM> further includes one or more cameras (not shown) in the treatment room for patient monitoring.

<FIG> schematically illustrates a drive stand <NUM> and gantry <NUM> of RT system <NUM>, according to various embodiments of the current disclosure. Covers, base positioning assembly <NUM>, couch <NUM>, and other components of RT system <NUM> are omitted in <FIG> for clarity. Drive stand <NUM> is a fixed support structure for components of RT treatment system <NUM>, including gantry <NUM> and a drive system <NUM> for rotatably moving gantry <NUM>. Drive stand <NUM> rests on and/or is fixed to a support surface that is external to RT treatment system <NUM>, such as a floor of an RT treatment facility. Gantry <NUM> is rotationally coupled to drive stand <NUM> and is a support structure on which various components of RT system <NUM> are mounted, including a linear accelerator (LINAC) <NUM>, an MV electronic portal imaging device (EPID) <NUM>, an imaging X-ray source <NUM>, and an X-ray imager <NUM>. During operation of RT treatment system <NUM>, gantry <NUM> rotates about bore <NUM> when actuated by drive system <NUM>.

Drive system <NUM> rotationally actuates gantry <NUM>. In some embodiments, drive system <NUM> includes a linear motor that can be fixed to drive stand <NUM> and interacts with a magnetic track (not shown) mounted on gantry <NUM>. In other embodiments, drive system <NUM> includes another suitable drive mechanism for precisely rotating gantry <NUM> about bore <NUM>. LINAC <NUM> generates an MV treatment beam <NUM> of high energy X-rays (or in some embodiments electrons) and EPID <NUM> is configured to acquire X-ray images with treatment beam <NUM>. Imaging X-ray source <NUM> is configured to direct a conical beam of X-rays, referred to herein as imaging X-rays <NUM>, through an isocenter <NUM> of RT system <NUM> to X-ray imager <NUM>, and isocenter <NUM> typically corresponds to the location of a target volume <NUM> to be treated. In the embodiment illustrated in <FIG>, X-ray imager <NUM> is depicted as a planar device, whereas in other embodiments, X-ray imager <NUM> can have a curved configuration.

X-ray imager <NUM> receives imaging X-rays <NUM> and generates suitable projection images therefrom. According to certain embodiments, such projection images can then be employed to construct or update portions of imaging data for a digital volume that corresponds to a three-dimensional (3D) region that includes target volume <NUM>. That is, a 3D image of such a 3D region is reconstructed from the projection images. In the embodiments, cone-beam computed tomography (CBCT) and/or digital tomosynthesis (DTS) can be used to process the projection images generated by X-ray imager <NUM>. CBCT is typically employed to acquire projection images over a relatively long acquisition arc, for example over a rotation of <NUM>° or more of gantry <NUM>. As a result, a high-quality 3D reconstruction of the imaged volume can generated. CBCT is often employed at the beginning of a radiation therapy session to generate a set-up 3D reconstruction. For example, CBCT may be employed immediately prior to application of treatment beam <NUM> to generate a 3D reconstruction confirming that target volume <NUM> has not moved or changed shape.

In some embodiments, partial-data reconstruction may be performed by RT system <NUM> during portions of an IGRT process in which partial image data is employed to generate a 3D reconstruction of target volume <NUM>. For example, as treatment beam <NUM> is directed to isocenter <NUM> while gantry <NUM> rotates through a treatment arc, DTS image acquisitions can be performed to generate image data for target volume <NUM>. Because DTS image acquisition is performed over a relatively short acquisition arc, for example between about <NUM>° and <NUM>°, near real-time feedback for the shape and position of target volume <NUM> can be provided by DTS imaging during the IGRT process. Alternatively, CBCT may be employed during portions of an IGRT process to generate a 3D reconstruction of target volume <NUM>. According to various embodiments described below, higher framerate X-ray images having little or no increased image lag can be generated for either scenario. Such higher framerate X-ray images are highly beneficial for generating accurate image data for target volume <NUM>, either for CBCT or DTS image acquisition.

In some embodiments, X-ray imager <NUM> includes a glass plate with a matrix or array of pixel detector elements, or pixels, formed thereon that each convert incident X-ray photons to electrical charge. In embodiments in which X-ray imager <NUM> is configured as an indirect flat panel detector, a scintillator material in X-ray imager <NUM> 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 that is proportional to incident light intensity) for a different pixel of what will eventually become a digital image. An encoder included in X-ray imager <NUM> then interprets each of these voltages and assigns a value to each that is proportional to the voltage. One such embodiment of X-ray imager <NUM> is illustrated in <FIG>.

<FIG> schematically illustrates a cross-sectional view of X-ray imager <NUM>, according to one embodiment of the disclosure. As shown, X-ray imager <NUM> includes a photosensitive element and detector circuitry layer <NUM> formed on a substrate <NUM>. In addition, X-ray imager <NUM> includes a layer of scintillator material <NUM> formed on photosensitive element and detector circuitry layer <NUM>. Also shown are incident X-rays <NUM> that have passed through a patient, sample, or other object of interest after being generated by imaging X-ray source <NUM> (shown in <FIG>). Together, photosensitive element and detector circuitry layer <NUM>, substrate <NUM>, and scintillator material <NUM> form an X-ray imaging array <NUM>. It is noted that photosensitive element and detector circuitry layer <NUM> is generally formed from a plurality of processing layers, and that X-ray imaging array <NUM> may include additional material layers not illustrated in <FIG>.

Photosensitive element and detector circuitry layer <NUM> generally includes a plurality of pixel detector elements <NUM>. Each pixel detector element <NUM> includes a photosensitive element, such as a photodiode, a photogate, or a phototransistor, as well as any other circuitry suitable for operation as a pixel detector element in X-ray imager <NUM>. In some embodiments, the photosensitive elements of pixel detector element <NUM> are amorphous silicon-based semiconductor devices. Photosensitive element and detector circuitry layer <NUM> may also include thin-film transistors (TFTs) for reading out the digital signals from pixel detector elements <NUM>. Scintillator material <NUM> may include one or more material layers including, but no limited to, gadolinium oxisulfide (Gd2O2S:Tb), cadmium tungstate (CdWO4), bismuth germanate (Bi4Ge3O12 or BGO), cesium iodide (Csl), or cesium iodide thallium (Csl:TI)), among others.

In the embodiment illustrated in <FIG>, X-ray imager <NUM> is depicted as an indirect flat panel detector, in which X-ray photons are converted to other light photons that are in turn detected and converted into charge. In other embodiments, X-ray imager <NUM> can be a direct flat panel detector (FPD). In a direct FPD, incident X-ray photons are converted directly into charge in an amorphous selenium layer, and the resultant charge pattern therein is read out by suitable hardware, such as a thin-film transistor (TFT) array, an active matrix array, microplasma line addressing, or the like.

In the embodiment illustrated in <FIG>, RT system <NUM> includes a single X-ray imager and a single corresponding imaging X-ray source. In other embodiments, RT system <NUM> can include two or more X-ray imagers, each with a corresponding imaging X-ray source. Thus, in such embodiments, RT system <NUM> includes a first imaging X-ray source and a corresponding X-ray imager mounted on gantry <NUM> and a second imaging X-ray source and corresponding X-ray imager mounted on gantry <NUM>. In such embodiments, the inclusion of multiple X-ray imagers in RT system <NUM> facilitates the generation of projection images (for reconstructing the target volume) over a shorter image acquisition arc. For instance, when RT system <NUM> includes two X-ray imagers and corresponding X-ray sources, an image acquisition arc for acquiring projection images of a certain image quality can be approximately half that for acquiring projection images of a similar image quality with a single X-ray imager and X-ray source. Further, in such embodiments, the inclusion of multiple X-ray imagers in RT system <NUM> facilitates the use of multiple X-ray source energies, since the first imaging X-ray source and the second imaging X-ray source can each operate at a different energy. Alternatively, in embodiments in which RT system <NUM> includes a single imaging X-ray source, the single imaging X-ray source can be configured as a multi-energy source.

The projection images generated by X-ray imager <NUM> are used to construct imaging data for a digital volume of patient anatomy within a 3D region that includes the target volume. Alternatively or additionally, such projection images can be used to update portions of existing imaging data for the digital volume corresponding to the 3D region. One embodiment of such a digital volume is described below in conjunction with <FIG>.

<FIG> schematically illustrates a digital volume <NUM> that is constructed based on projection images generated by one or more X-ray imagers included in RT system <NUM>, according to various embodiments of the current disclosure. For example, in some embodiments, the projection images can be generated by a single X-ray imager, such as X-ray imager <NUM>, and in other embodiments the projection images can be generated by multiple X-ray imagers.

Digital volume <NUM> includes a plurality of voxels <NUM> (dashed lines) of anatomical image data, where each voxel <NUM> corresponds to a different location within digital volume <NUM>. For clarity, only a single voxel <NUM> is shown in <FIG>. Digital volume <NUM> corresponds to a 3D region that includes target volume <NUM>. In <FIG>, digital volume <NUM> is depicted as an 8x8x8 voxel cube, but in practice, digital volume <NUM> generally includes many more voxels, for example orders of magnitude more than are shown in <FIG>.

For purposes of discussion, target volume <NUM> can refer to the gross tumor volume (GTV), clinical target volume (CTV), or the planning target volume (PTV) for a particular treatment. The GTV depicts the position and extent of the gross tumor, for example what can be seen or imaged; the CTV includes the GTV and an additional margin for sub-clinical disease spread, which is generally not imageable; and the PTV is a geometric concept designed to ensure that a suitable radiotherapy dose is actually delivered to the CTV without adversely affecting nearby organs at risk. Thus, the PTV is generally larger than the CTV, but in some situations can also be reduced in some portions to provide a safety margin around an organ at risk. The PTV is typically determined based on imaging performed prior to the time of treatment, and alignment of the PTV with the current position of patient anatomy at the time of treatment is facilitated by embodiments of the disclosure.

According to various embodiments described below, image information associated with each voxel <NUM> of digital volume <NUM> is constructed from projection images generated by single or multiple X-ray imagers, for example via a CBCT or DTS process. In some embodiments, image information associated with some or all of voxels <NUM> of digital volume <NUM> is updated via projection images generated by the 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 completed. In this way, the location and shape of target volume <NUM> can be confirmed while the treatment is underway. Thus, if a sufficient portion of the target volume <NUM> is detected to extend outside a threshold region, the treatment can either be aborted or modified. In such an instance, modification of the treatment can be accomplished by adjusting patient position and/or the treatment beam.

<FIG> is a partial circuit diagram <NUM> of photosensitive element and detector circuitry layer <NUM>, according to one embodiment of the disclosure. Photosensitive element and detector circuitry layer <NUM> can be included in a suitable X-ray detector panel, such as X-ray imager <NUM> in <FIG>. As shown, photosensitive element and detector circuitry layer <NUM> includes a plurality of pixel detector elements <NUM> that are each communicatively coupled to a readout stage <NUM> via a common data line <NUM>. Generally, photosensitive element and detector circuitry layer <NUM> includes an M x N matrix of pixel detector elements <NUM>, where M equals the number of rows of pixel detector elements <NUM> and N equals the number of columns of pixel detector elements <NUM>. Generally, M and N have values on the order of about <NUM> to <NUM> in an X-ray detector panel. For clarity, only a single array of pixel detector elements <NUM> that forms one column of M pixel detector elements <NUM> is shown in <FIG>. In practice, photosensitive element and detector circuitry layer <NUM> includes a total of N such arrays of pixel detector elements <NUM>, where each array forms one of the M rows of pixels of an X-ray detector panel.

In the embodiment illustrated in <FIG>, each pixel detector element <NUM> includes a photodiode <NUM> and a readout switch <NUM> that communicatively couples the photodiode <NUM> to readout stage <NUM>. As shown, each photodiode <NUM> is communicatively coupled to a bias voltage Vbias and is also communicatively coupled to data line <NUM> via a respective readout switch <NUM>. Each readout switch <NUM> is typically formed as part of the associated pixel detector element <NUM>. Alternatively, in some embodiments, each readout switch <NUM> is formed proximate to the associated pixel detector element <NUM>. In either case, readout switches <NUM> are generally formed as part of photosensitive element and detector circuitry layer <NUM>. For example, in some embodiments, readout switches <NUM> are implemented as thin-film transistors (TFTs) that are formed on the same substrate as the pixel detector elements <NUM>.

Readout stage <NUM> is a readout device configured to readout accumulated charge from the pixel detector elements <NUM> of a particular column <NUM> of pixel detector elements <NUM>. Readout stage <NUM> reads out a particular pixel detector element <NUM> (for example, the pixel detector element <NUM> associated with Pixel M-<NUM>) when the readout switch <NUM> for that particular pixel detector element <NUM> (for example, readout switch <NUM> associated with Pixel M-<NUM>) closes and communicatively couples that pixel detector element <NUM> to data line <NUM>. In operation, for each other column (not shown) of photosensitive element and detector circuitry layer <NUM>, a single pixel detector element <NUM> can be simultaneously readout by the readout stage <NUM> associated with the column. Thus, a complete row of pixel detector elements <NUM> of photosensitive element and detector circuitry layer <NUM> can be read out at one time by the appropriately timed closing of one readout switch <NUM> in each column of pixel detector elements <NUM>.

Readout stage <NUM> is configured to convert analog signals, such as charge accumulated in pixel detector elements <NUM>, to digital X-ray image signals. In some embodiments, readout stage <NUM> includes conversion circuitry <NUM> for converting such signals to digital X-ray image signals. Conversion circuitry <NUM> can include any technically feasible circuitry suitable for performing such conversions. For example, in some embodiments, conversion circuitry <NUM> includes an analog-to-digital converter, an analog front-end, or the like. In addition, in the embodiment illustrated in <FIG>, readout stage <NUM> includes a reset switch <NUM> that is configured to communicatively couple data line <NUM> to a reference voltage Vref, such as ground or any other suitable reference voltage. Thus, when a readout switch <NUM> of a particular pixel detector element <NUM> is closed while reset switch <NUM> is closed, charge currently accumulated in that particular pixel detector element <NUM> discharges down to the reference voltage Vref. By contrast, when a readout switch <NUM> of a particular pixel detector element <NUM> is closed while reset switch <NUM> is open, charge currently accumulated in that particular pixel detector element <NUM> discharges to the readout stage <NUM> and is read by conversion circuitry <NUM>.

In the embodiment illustrated in <FIG>, reset switch <NUM> is included in readout stage <NUM>. In other embodiments, a reset switch is still communicatively coupled to data line <NUM>, but is implemented outside of readout stage <NUM>. One such embodiment is illustrated in <FIG> is a partial circuit diagram <NUM> of photosensitive element and detector circuitry layer <NUM>, according to another embodiment of the disclosure. As shown, in the embodiment illustrated in <FIG>, a reset switch <NUM> is configured to selectively couple data line <NUM> to reference voltage Vref, but is located at some other position along data line <NUM> than readout stage <NUM>. In yet other embodiments, a reset switch is communicatively coupled to data line <NUM>, but is implemented via existing circuitry included in conversion circuitry <NUM>. One such embodiment is illustrated in <FIG> is a partial circuit diagram <NUM> of photosensitive element and detector circuitry layer <NUM>, according to yet another embodiment of the disclosure. As shown, in the embodiment illustrated in <FIG>, a reset switch <NUM> is configured to selectively couple data line <NUM> to reference voltage Vref and is included within conversion circuitry <NUM>. Alternatively, the functionality of reset switch <NUM> is implemented by one or more components of conversion circuitry <NUM>, and reset switch <NUM> is not a dedicated switch or transistor for coupling data line <NUM> to reference voltage Vref.

According to various embodiments described herein, the acquisition of an X-ray image with an X-ray imager, such as X-ray imager <NUM>, is performed in three phases: an irradiation phase, a readout phase, and a reset phase. In the irradiation phase, each pixel detector element <NUM> integrates the charge that is generated through irradiation of the panel with imaging X-rays, via pixel capacitance. In the readout phase, the accumulated charge of each pixel detector element <NUM> is transferred to readout stage <NUM> and is processed. In the reset phase, residual charge is transferred from each pixel detector element <NUM>, thereby minimizing or otherwise reducing image lag present in the next X-ray image to be acquired. One such embodiment is described below in conjunction with <FIG>.

<FIG> is a timing diagram <NUM> schematically illustrating charge accumulation and loss in pixel detector elements <NUM> during an irradiation phase <NUM>, a readout phase <NUM>, and a reset phase <NUM> of a single X-ray image acquisition, according to an embodiment of the present disclosure. More specifically, charge accumulation and loss is shown for M rows of pixel detector elements <NUM> in an X-ray detector panel that includes an M x N matrix of pixel detector elements <NUM>. One such row is described above in conjunction with <FIG>.

As shown, throughout irradiation phase <NUM>, each of the M rows are simultaneously irradiated, and charge is accumulated in the pixel detector elements <NUM> of each row. In readout phase <NUM>, charge accumulated in the pixel detector elements <NUM> of each row are read out sequentially. That is, the pixel detector elements <NUM> for row <NUM> are read out by readout stage <NUM>, then the pixel detector elements <NUM> for row <NUM> are read out by readout stage <NUM>, and so on, until the pixel detector elements <NUM> of all M rows are read out and an X-ray image can be generated. In reset phase <NUM>, residual charge that remains in the pixel detector elements <NUM> of each row is concurrently transferred from all rows of an X-ray detector panel. Readout phase <NUM> is described in greater detail below in conjunction with <FIG>, and reset phase <NUM> is described in greater detail below in conjunction with <FIG>.

<FIG> is a schematic timing diagram <NUM> illustrating charge loss for two adjacent rows of pixel detector elements <NUM> during readout phase <NUM>, according to an embodiment of the present disclosure. In <FIG>, a portion of readout phase <NUM> is shown for a first array (row M-<NUM>) of pixel detector elements <NUM> and an adjacent second array (row M-<NUM>) of pixel detector elements <NUM>. During readout phase <NUM>, accumulated charge in each of the pixel detector elements <NUM> of row M-<NUM> is read during a first readout time interval tread1, and accumulated charge in each of the pixel detector elements <NUM> of row M-<NUM> is read during a second readout time interval tread2, that follows first readout time interval tread1.

First readout time interval tread1 begins after the accumulated charge in each of the pixel detector elements <NUM> of the preceding row (e.g., row M-<NUM>, not shown) has been read out, since photosensitive element and detector circuitry layer <NUM> is typically connected to a readout stage <NUM> that is configured to read out one row of pixel detector elements <NUM> at a time. Similarly, second readout time interval tread2 begins after the accumulated charge in each of the pixel detector elements <NUM> of row M-<NUM> has been read out and first readout time interval tread1 has ended.

During first readout time interval tread1, the readout switch <NUM> for each pixel detector element <NUM> of row M-<NUM> closes (indicated by ON state in <FIG>) and an accumulated charge QM-<NUM> decreases in magnitude over first readout time interval tread1 from an initial charge value Q<NUM>M-<NUM> to a remainder charge value QRemM-<NUM>. The readout switch <NUM> for each pixel detector element <NUM> of row M-<NUM> then opens (indicated by OFF state in <FIG>) and no more charge is read out from the pixel detector element <NUM> of row M-<NUM>. Typically, the value of remainder charge value QRemM-<NUM> is a function of the duration of first readout time interval tread1, a time dependent release of trapped charges, and a pixel time constant τpix, which is substantially the same for each photodiode <NUM> of an X-ray imager. For example, in some embodiments, the value of remainder charge value QRemM-<NUM> can be determined based on Equation <NUM>:
<MAT>.

Similarly, during second readout time interval tread2, the readout switch <NUM> for each pixel detector element <NUM> of row M-<NUM> closes (indicated by ON state in <FIG>) and an accumulated charge QM-<NUM> decreases in magnitude over second readout time interval tread2 from an initial charge value Q0M-<NUM> to a remainder charge value QRemM-<NUM>. The readout switch <NUM> for each pixel detector element <NUM> of row M-<NUM> then opens (indicated by OFF state in <FIG>) and no more charge is read out from the pixel detector element <NUM> of row M-<NUM>. This process continues sequentially through the remaining rows of the X-ray imager.

Charge cannot escape from a photodiode <NUM> once the associated readout switch <NUM> opens. Therefore, in a conventional X-ray imager, the value of remainder charge value QRem for a row corresponds to image lag for each of the pixel detector element <NUM> of the row, since such remaining charge is present when the subsequent irradiation phase <NUM> begins. As a result, image quality suffers. Alternatively, the duration of each readout time interval can be increased so that the magnitude of remainder charge value QRem is irrelevant. In the latter case, panel readout time of the X-ray imager is greatly slowed, since the panel readout time increases based on the relation: (readout time interval increase) x (number of rows of pixel detector elements). Further, because the rate at which remainder charge value QRem decays during readout is an exponential function, a relatively large increase in the readout time interval is required to produce even a small reduction in the remainder charge value QRem. By contrast, according to various embodiments described herein, reset phase <NUM> enables remainder charge value QRem for each pixel of an X-ray detector panel to be greatly reduced over a relatively short time interval prior to the next irradiation phase <NUM>. Consequently, image lag can be prevented without slowing panel readout time by more than the duration of reset phase <NUM>.

<FIG> is a schematic timing diagram <NUM> illustrating charge loss for two adjacent rows of pixel detector elements <NUM> during reset phase <NUM>, according to an embodiment of the present disclosure. In <FIG>, a portion of reset phase <NUM> is shown for a first array (row M-<NUM>) of pixel detector elements <NUM> and an adjacent second array (row M-<NUM>) of pixel detector elements <NUM>. Reset phase <NUM> begins after readout phase <NUM> for all M rows of the X-ray detector panel has been completed.

During reset phase <NUM>, residual charge in each of the pixel detector elements <NUM> is concurrently transferred from each of the M rows of the X-ray detector panel, for example during a single reset time interval treset. Thus, during reset time interval treset, residual charge in each of the pixel detector elements <NUM> of row M-<NUM> is transferred from the pixel detector elements, for example to ground or some other suitable reference voltage Vref. In addition, during reset time interval treset, residual charge in each of the pixel detector elements <NUM> of row M-<NUM> is transferred from the pixel detector elements to reference voltage Vref. Further, during reset time interval treset, residual charge in each of the pixel detector elements <NUM> of the remaining rows (not shown) of the X-ray detector panel is transferred from the pixel detector elements to reference voltage Vref. Thus, during reset time interval treset, residual charge from some or all of the M rows is concurrently transferred from pixel detector elements <NUM> to reference voltage Vref. As a result, the magnitude of image lag associated with each pixel detector element <NUM> can be reduced much more quickly than by increasing the duration of the readout time interval tread for each row of pixel detector elements <NUM>.

As shown in <FIG>, at the beginning of reset time interval treset, residual charge in each particular pixel detector element <NUM> of row M-<NUM> equals the remainder charge value QRemM-<NUM>, and at the end of reset time interval treset, residual charge in each particular pixel detector element <NUM> of row M-<NUM> equals a final charge value QFinM-<NUM> for that particular pixel detector element <NUM>. Similarly, at the beginning of reset time interval treset, residual charge in each particular pixel detector element <NUM> of row M-<NUM> equals the remainder charge value QRemM-<NUM> of that particular pixel detector element <NUM>, and at the end of reset time interval treset, residual charge in each particular pixel detector element <NUM> of row M-<NUM> equals a final charge value QFin M-<NUM> for that particular pixel detector element <NUM>. Similar to remainder charge value QRem, in some embodiments, a final charge value QFin for each pixel detector element <NUM> of a particular row (for example row M-<NUM>) can be determined based on Equation <NUM>:
<MAT>.

Thus, the magnitude of QFin for pixel detector elements <NUM> decays at substantially the same rate during reset time interval treset of reset phase <NUM> as the magnitude of QRem decays during one of the readout time intervals of readout phase <NUM>, such as readout time interval tread1. However, in reset phase <NUM>, charge is transferred from some, most, or all of the M rows of the pixel detector elements <NUM> of the X-ray detector panel simultaneously. Thus, increasing the duration of reset time interval treset can greatly reduce the magnitude of accumulated charge (from QRem to QFin) in each of the pixel detector elements <NUM> of an X-ray detector panel without significantly increasing the panel readout time for the X-ray detector panel. By contrast, to achieve the same reduction in accumulated charge (down to QFin) by increasing the duration of each of the readout time intervals of readout phase <NUM>, the panel readout time for the X-ray detector panel is greatly increased. For example, in one embodiment, reset time interval treset of reset phase <NUM> is set to be three times as long as the readout time intervals of readout phase <NUM>. Therefore, panel readout time is increased by <NUM> x (readout time interval treset). To achieve the same reduction in accumulated charge in the pixel detector elements <NUM> of the X-ray detector panel without using the reset phase <NUM>, the panel readout time is quadrupled, which significantly slows imaging framerate of the X-ray detector.

<FIG> sets forth a flowchart 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. The method may include one or more operations, functions, or actions as illustrated by one or more of blocks <NUM>-<NUM>. Although the blocks are illustrated in a sequential order, these blocks may be performed in parallel, and/or in a different order than those described herein. Also, the various blocks may be combined into fewer blocks, divided into additional blocks, and/or eliminated based upon the desired implementation. The control algorithms for the method steps can be implemented in whole or in part as software- or firmware-implemented logic, and/or as hardware-implemented logic circuits. Further, the control algorithms for the method steps can be performed in whole or in part by treatment control computer <NUM> of RT system <NUM> (shown in <FIG>), a controller included in X-ray imager <NUM> (shown in <FIG>), any other suitable controller associated with RT system <NUM>, or any combination thereof.

A method <NUM> begins at step <NUM>, where treatment control computer <NUM>, a controller included in X-ray imager <NUM>, or any other suitable controller associated with RT system <NUM> causes X-ray imager <NUM> to be prepared for irradiation phase <NUM> and the acquisition of an X-ray image. In some embodiments, the controller causes charge currently accumulated in pixel detector elements <NUM> of X-ray imager <NUM> to be transferred to ground or other reference voltage Vref. For example, the controller causes readout switch <NUM> of each pixel detector element <NUM> and reset switch <NUM> to close for a certain time interval to reduce accumulated charge in pixel detector elements <NUM>. The controller then causes readout switch <NUM> of each pixel detector element <NUM> and reset switch <NUM> to open.

In step <NUM>, the controller begins irradiation phase <NUM>. For example, in some embodiments, the controller causes imaging X-ray source <NUM> to direct imaging X-rays <NUM> through isocenter <NUM> of RT system <NUM> to X-ray imager <NUM>.

In step <NUM>, X-ray imager <NUM> receives imaging X-rays <NUM> and charge is accumulated in some or all of pixel detector elements <NUM>.

In step <NUM>, the controller ends irradiation phase <NUM>. For example, in some embodiments, the controller causes imaging X-ray source <NUM> to stop directing imaging X-rays <NUM> to X-ray imager <NUM>.

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

In step <NUM>, the controller causes accumulated charge from the pixel detector elements <NUM> of the selected pixel array to be read out during a readout interval. For example, in some embodiments, the controller causes the readout switch <NUM> of each pixel detector element <NUM> in the selected pixel array to close for a readout time interval tread while reset switch <NUM> to reference voltage Vref remains open. As a result, charge currently accumulated in each pixel detector element <NUM> in the selected pixel array discharges to the readout stage <NUM>, is read by conversion circuitry <NUM>, and is processed as part of the current X-ray image being acquired. During implementation of step <NUM>, accumulated charge present in each pixel detector element <NUM> in the selected pixel array is reduced from an initial charge value Q<NUM> to a remainder charge value QRem at a rate described by previously presented Equation <NUM>. It is noted that initial charge value Q<NUM> is generally different for each pixel detector element <NUM>. Similarly, remainder charge value QRem is a function of initial charge value Q<NUM> and therefore is also generally different for each pixel detector element <NUM>.

In step <NUM>, the controller determines whether there are remaining pixel arrays to be read out. If yes, method <NUM> returns to step <NUM>; if no, method <NUM> proceeds to step <NUM>.

In step <NUM>, the controller begins reset phase <NUM>. For example, in some embodiments, the controller causes reset switch <NUM> to close, and data line <NUM> is communicatively coupled to Vref.

In step <NUM>, the controller selects one or more pixel arrays of X-ray detector <NUM> to be reset. For example, in the embodiment of reset phase <NUM> illustrated in <FIG>, the controller selects multiple pixel arrays of X-ray detector <NUM>, i.e., all M rows of pixel detector elements <NUM>. In such embodiments, remainder charge value QRem for the pixel detector elements <NUM> of all M rows is reduced simultaneously. Alternatively, in some embodiments, a staged reset of pixel arrays is performed in a reset phase. In such embodiments, the peak current that is exposed to readout stage <NUM> during reset phase <NUM> can be reduced by preventing all pixel arrays of X-ray detector <NUM> from simultaneously being coupled to data line <NUM>. Thus, in some embodiments, in step <NUM>, the controller selects a portion of the total M rows of pixel arrays to begin being reset in step <NUM>, for example one tenth of the M rows. In such embodiments, subsequent portions are selected to begin being reset while the previously selected portions continue to be reset. One such embodiment is illustrated in <FIG> and <FIG>.

<FIG> is a timing diagram <NUM> schematically illustrating charge loss for different rows of pixel detector elements <NUM> during reset phase <NUM>, according to an embodiment of the present disclosure. <FIG> illustrates a portion of reset phase <NUM> for a first array (row M-<NUM>) of pixel detector elements <NUM> and an adjacent second array (row M-<NUM>) of pixel detector elements <NUM>, according to an embodiment of the present disclosure.

Reset phase <NUM> is similar to reset phase <NUM> of <FIG>, in that during reset phase <NUM>, residual charge in each pixel detector element <NUM> of multiple rows of an X-ray detector panel is transferred from the pixel detector elements concurrently. That is, during at least a portion of reset phase <NUM>, multiple rows of an X-ray detector panel are undergoing reset simultaneously and accumulated charge is being transferred simultaneously from all pixel detector elements <NUM> associated with the multiple rows. However, in reset phase <NUM>, the reset time interval treset begins at different times for different rows of pixel detector elements <NUM>. Thus, reset of the M rows of pixel detector elements <NUM> is staged throughout reset phase <NUM>. Consequently, the larger initial current in data line <NUM> that results when a row of pixel detector elements <NUM> is initially coupled thereto does not occur for all rows of pixel detector elements <NUM> simultaneously. Instead, the rows of pixel detector elements <NUM> are initially coupled to data line <NUM> and readout stage <NUM> at different times in reset phase <NUM>.

In the embodiment illustrated in <FIG> and <FIG>, each of the M rows of pixel detector elements <NUM> is initially coupled to readout stage <NUM> at a different time in reset phase <NUM>. Thus, in such an embodiment, each row of pixel detector elements <NUM> is associated with a different reset initiating time. For example, as shown in <FIG>, reset of the pixel detector elements <NUM> of row M-<NUM> is initiated at a reset initiating time trese,M-<NUM>, while reset of the pixel detector elements <NUM> of row M-<NUM> is initiated at a reset initiating time treset,M-<NUM>. In some embodiments, treset,M-<NUM> is separated in time from reset initiating time treset,M-<NUM> by a staging time interval tstage. In some embodiments, staging time interval tstage is generally uniform between consecutive rows of pixel detector elements <NUM>, while in other embodiments, staging time interval tstage can vary between different consecutive rows of pixel detector elements <NUM>. Generally, the duration of staging time interval tstage is relatively small compared to the duration of reset phase <NUM>. Consequently, in embodiments in which the reset of the M rows of pixel detector elements <NUM> is staged throughout reset phase <NUM>, the overall duration of reset phase <NUM> is not significantly increased.

In the embodiment illustrated in <FIG> and <FIG>, each of the M rows of pixel detector elements <NUM> is initially coupled to readout stage <NUM> at a unique time in reset phase <NUM>. Alternatively, in some embodiments, a first group of multiple rows of pixel detector elements <NUM> are initially coupled to readout stage <NUM> at one time in reset phase <NUM>, a second group of multiple rows of pixel detector elements <NUM> are initially coupled to readout stage <NUM> at a later time in reset phase <NUM>, a third group of multiple rows of pixel detector elements <NUM> are initially coupled to readout stage <NUM> at another later time in reset phase <NUM>, and so on. For example, in one such embodiment, each such group of multiple rows of pixel detector elements <NUM> includes <NUM>% of the M rows of pixel detector elements <NUM> and each such group includes different rows of pixel detector elements <NUM> than the other rows of pixel detector elements. Thus, in such an embodiment, reset phase <NUM> includes <NUM> different reset initiating times, rather than one reset initiating time for each row of pixel detector elements.

Returning to <FIG>, in step <NUM>, the controller causes the row or rows of pixel detector elements <NUM> selected in step <NUM> to be reset. For example, in some embodiments, the controller causes reset switch <NUM> to close, thereby communicatively coupling data line <NUM> to reference voltage Vref. Furthermore, the controller causes the readout switch <NUM> of each pixel detector element <NUM> in the selected row or rows to close, thereby communicatively coupling such pixel detector elements <NUM> to data line <NUM>. As a result, charge present in the pixel detector elements <NUM> (e.g., remainder charge value QRem) begins to be transferred to reference voltage Vref. Thus, the charge present in the pixel detector elements <NUM> is reduced at a rate indicated by previously presented Equation <NUM>.

In embodiments in which all M rows of pixel detector elements <NUM> are selected in step <NUM>, all pixel detector elements <NUM> of X-ray imager <NUM> are reset simultaneously. In such embodiments, after reset time interval treset has elapsed, method <NUM> proceeds to step <NUM>.

In embodiments in which a staged reset of pixel arrays is performed in the reset phase, a single row of pixel detector elements <NUM> begins to be reset in step <NUM> or a group of rows of pixel detector elements <NUM> begins to be reset in step <NUM>. In such embodiments, after a staging time interval tstage has elapsed, method <NUM> proceeds to step <NUM>.

In step <NUM>, the controller determines whether there are remaining pixel arrays that have not yet begun being reset. If yes, method returns to step <NUM>; if no, method <NUM> proceeds to step <NUM>. It is noted that in embodiments in which a staged reset of pixel arrays is performed in the reset phase, method <NUM> may return to step <NUM> while some pixel arrays of X-ray imager <NUM> are currently being reset. That is, the controller may perform step <NUM> before reset time interval treset has elapsed for some or all pixel arrays that have been selected to be reset.

In step <NUM>, the controller determines whether reset has been completed for all pixel arrays. That is, the controller determines whether reset time interval treset has elapsed for all pixel arrays of X-ray imager <NUM>. If no, method <NUM> returns back to step <NUM>; if yes, method <NUM> proceeds to step <NUM>.

In step <NUM>, the controller ends the reset phase. For example, in some embodiments, the controller causes reset switch <NUM> to open and the readout switch <NUM> of each pixel detector element <NUM> of X-ray imager <NUM> to open. When acquisition of further X-ray images is planned, method <NUM> returns back to step <NUM>.

Implementation of method <NUM> enables a significant reduction in charge remaining in pixel detector elements <NUM> (for example from QRem to QFin) over a reset time interval treset. As noted above, reset time interval treset can be selected to be relatively short compared to the duration of a readout time interval. In a conventional X-ray imager, an equivalent charge reduction can only be realized over a much greater time interval, i.e., (reset time interval treset) x (number of rows of pixel detector elements <NUM>). Since the number of rows of pixel detector elements <NUM> can be on the order of <NUM> to <NUM>, the panel readout time for the conventional X-ray imager can be significantly increased, preventing fast image acquisition.

In practice, a treatment beam in an RT system typically generates a large amount of scattered radiation in all directions, including that emanating from the patient, treatment table, and machine components. As a result, a large amount of MV scatter can be incident on an X-ray imager (e.g., X-ray imager <NUM> in <FIG>). In some instances, the amount of such X-ray scatter can even exceed the magnitude of imaging X-rays. Accordingly, in some embodiments, a reset phase, such as reset phase <NUM> or reset phase <NUM>, is timed to coincide with the application of a treatment beam, such as treatment beam <NUM>. X-ray imager <NUM> is insensitive to radiation during a reset phase as described herein. That is, pixel detector elements <NUM> of X-ray imager <NUM> do not accumulate charge during such a reset phase, even when treatment beam <NUM> produces significant X-ray scatter that is incident on X-ray imager <NUM>. As a result, the implementation of the reset phase to coincide with a burst or other application of treatment beam <NUM> reduces noise that is normally induced by treatment beam <NUM>. One such embodiment is illustrated in <FIG>.

<FIG> is a timing diagram <NUM> schematically illustrating the relative timing of an irradiation phase <NUM>, a readout phase <NUM>, and a reset phase <NUM> relative to the delivery of a treatment beam pulse <NUM>, according to an embodiment of the present disclosure. As shown, a KV (imaging) beam pulse <NUM> occurs during each irradiation phase <NUM> and delivery of treatment beam pulse <NUM> is timed to occur during each reset phase <NUM>. As described above, pixel detector elements <NUM> of X-ray imager <NUM> do not accumulate charge during a reset phase. Consequently, even though scattered X-rays from treatment beam pulses <NUM> may strike the scintillator layer of X-ray imager <NUM>, noise is not added to the image data generated by X-ray imager <NUM>.

While the 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 can also be implemented in a hand-held or portable flat panel X-ray detector (FPD), a statically mounted FPD, an X-ray imager configured for dynamic X-ray imaging, such as a fluoroscopic imaging, and the like. Further, the embodiments can be employed for the detection in intra-fraction motion or pre-treatment imaging prior to treatment.

Aspects of the present embodiments may be embodied as a system, method or computer program product. Accordingly, aspects of the present disclosure may take the form of an entirely hardware embodiment, an entirely software embodiment (including firmware, resident software, micro-code, etc.) or an embodiment combining software and hardware aspects that may all generally be referred to herein as a "circuit," "module" or "system. " Furthermore, aspects of the present disclosure may take the form of a computer program product embodied in one or more computer readable medium(s) having computer readable program code embodied thereon.

More specific examples (a nonexhaustive list) of the computer readable storage medium would include the following: an electrical connection having one or more wires, a portable computer diskette, 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 disc read-only memory (CD-ROM), an optical storage device, a magnetic storage device, or any suitable combination of the foregoing.

Claim 1:
A method of acquiring an X-ray image of a target volume (<NUM>) for an image guided radiation therapy, IGRT, system configured to apply a treatment beam (<NUM>, <NUM>), the method comprising:
transferring accumulated charge from each pixel (<NUM>) in a first array of pixels of an X-ray detector panel (<NUM>) to a readout device (<NUM>) during a first readout interval, wherein residual charge (QRem) remains in each pixel (<NUM>) in the first array of pixels after the transferring, the residual charge (QRem) corresponding to image lag of the pixel (<NUM>);
after the first readout interval, transferring accumulated charge from each pixel (<NUM>) in a second array of pixels of the X-ray detector panel (<NUM>) to the readout device (<NUM>) during a second readout interval, wherein residual charge remains in each pixel (<NUM>) in the second array of pixels after the transferring of accumulated charge from each pixel (<NUM>) in the second array of pixels, the residual charge (QRem) corresponding to image lag of the pixel (<NUM>); and
during a reset interval (<NUM>) that follows the second readout interval, concurrently transferring at least a portion of the residual charge from each pixel (<NUM>) in the first array of pixels and at least a portion of the residual charge from each pixel (<NUM>) in the second array of pixels to the reduce image lag associated with the pixels (<NUM>);
wherein a start time of the reset interval (<NUM>) is selected to occur prior to application of the treatment beam (<NUM>, <NUM>) and an end time of the reset interval (<NUM>) is selected to occur after the application of the treatment beam (<NUM>, <NUM>).