Patent Description:
Radiation therapy is a localized treatment for a specific target tissue (a planning target volume), such as a cancerous tumor. Ideally, radiation therapy is performed on the planning target volume that spares the surrounding normal tissue from receiving doses above specified tolerances, thereby minimizing risk of damage to healthy tissue. Prior to the delivery of radiation therapy, an imaging system is typically employed to provide a three dimensional image of the target tissue and surrounding area. From such imaging, the size and mass of the target tissue can be estimated and an appropriate treatment plan generated and planning target volume determined.

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 while radiation treatment is delivered 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. In this context, reference is made to documents <CIT> and <CIT>.

In accordance with at least some embodiments of the present disclosure, a radiation therapy system is configured to generate volumetric image data for a target volume during treatment, where the volumetric image data are not degraded by X-ray image noise caused by the scatter of treatment X-rays. As a result, the volumetric image data so generated can be employed to more accurately detect intra-fraction motion that occurs during application of treatment X-rays. For example, with such higher quality volumetric image data of a target volume, intra-fraction motion of the target volume caused by loss of breath-hold and/or anatomical variations due to peristalsis can be more readily detected. Thus, the radiation therapy system can perform image guided radiation therapy (IGRT) that monitors intra-fraction motion using X-ray imaging rather than magnetic resonance imaging (MRI). Detected anatomical variations can then either be compensated for, via patient repositioning and/or treatment modification, or the current treatment can be aborted.

In some embodiments, time-domain interleaving of treatment X-rays and imaging X-rays during an IGRT process prevents scatter of the treatment X-rays from degrading the quality of X-ray images used to generate volumetric image data of a target volume. In such embodiments, imaging X-rays are delivered to the target volume during one or more imaging intervals, and one or more pulses of treatment X-rays are delivered to the target volume between the imaging intervals. In instances in which a pulse of treatment X-rays is timed to occur during an imaging interval, the pulse of treatment X-rays is inhibited from occurring during the imaging interval and is rescheduled to occur at a later time that does not coincide with that imaging interval or subsequent imaging intervals.

In some embodiments, treatment X-rays are delivered to a target volume at the same time that imaging X-rays are also delivered to the target volume for generating volumetric image data of the target volume. That is, during an imaging interval in which imaging X-rays are delivered to the target volume, one or more pulses of treatment X-rays are also delivered to the target volume. In such embodiments, in each pixel of an X-ray imaging device, image signal is accumulated during portions of the imaging interval in which only treatment X-rays are delivered to the target volume and is prevented from accumulating in each pixel during the pulses of treatment X-rays.

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. 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 scope of the subject matter presented here. 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 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-imaging device. Because there are time intervals in which treatment X-rays are applied to the target volume while imaging X-rays are received by the X-ray imaging device, the X-ray projection images generated during such time intervals can include significant image noise caused by scattered treatment X-rays that are captured by the X-ray imaging device. For example, when typical megavolt (MV) treatment X-rays and kilovolt (kV) imaging X-rays are employed, the magnitude of scattered MV radiation from the patient, treatment table, and machine components can exceed all other image noise. Because the Poisson-distributed component of such image noise cannot be filtered, the image quality of the X-ray projection images is reduced. Consequently, detection of intra-fraction motion based on such X-ray images is negatively impacted.

In light of the above, there is a need in the art for improved systems and techniques for ensuring a target volume remains properly positioned for treatment in a radiation therapy system while a treatment beam is delivered to the target volume. According to various embodiments described herein, a radiation system is configured to generate high-quality X-ray images of the target volume that are not degraded by image noise from scattered treatment X-rays. 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.

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.

By contrast, partial-data reconstruction is 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 lGRT process.

In some embodiments, X-ray imager <NUM> includes a glass plate with a matrix or array of pixel detector elements 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 electric charge 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 signals and assigns a value to each that is proportional to the signal. 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>. For example, photosensitive element and detector circuitry layer <NUM> may also include thin-film transistors (TFTs) for reading out the digital signals from the pixel detector elements. 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:Tl)), among others.

In some embodiments, TFTs included in detector circuitry layer <NUM> include one or more specific semiconductor materials that enable incorporation of more transistors into a pixel detector element <NUM>, such as indium gallium zinc oxide (IGZO), a low-temperature polysilicon semiconductor material, and a polycrystalline silicon material. Alternatively, in some embodiments, transistors included in detector circuitry layer <NUM> can be based on complementary metal-oxide-semiconductor (CMOS) technology that enables more complex circuits to be included in a single pixel detector element <NUM>. Since the charge carrier mobility in CMOS is very high the transistors could be made very small and therefore charge injection, induced by switching, could be significantly reduced compared to other technologies.

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. One such embodiment is illustrated in <FIG>.

<FIG> schematically illustrates a drive stand <NUM> and gantry <NUM> of RT system <NUM>, according to various embodiments of the current disclosure. Drive stand <NUM> and gantry <NUM> are substantially similar in configuration to drive stand <NUM> and gantry <NUM> in <FIG>, except that the components of RT system <NUM> that are mounted on gantry <NUM> include a first imaging X-ray source <NUM>, a first X-ray imager <NUM>, a second imaging X-ray source <NUM>, and a second X-ray imager <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.

The projection images generated by X-ray imager <NUM> (or by first x-ray imager <NUM> and second 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, such as first x-ray imager <NUM> and second X-ray imager <NUM>.

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 imagable; 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 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.

During use, a treatment beam 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> or first X-ray imager <NUM> and second 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, time-domain interleaving of a treatment beam with imaging X-rays can be employed to reduce or eliminate interference from X-ray scatter of the treatment beam with the detection of imaging X-rays. That is, in such embodiments, timing of the delivery of imaging X-rays (e.g., imaging X-rays <NUM> in <FIG>) to target volume <NUM> and the treatment beam (e.g., treatment beam <NUM> in <FIG>) is optimized. In such embodiments, imaging X-rays and the treatment beam are pulsed or otherwise intermittently activated, so that when the imaging X-rays and are being received by an X-ray imager, the treatment beam is not being delivered to target volume <NUM>.

<FIG> is a circuit diagram <NUM> of pixel detector element <NUM>, according to an embodiment of the present disclosure. The embodiment of pixel detector element <NUM> illustrated by circuit diagram <NUM> is included in an X-ray imager that enables time-domain interleaving of a treatment beam with imaging X-rays. Pixel detector element <NUM> can be included in an X-ray imager of RT system <NUM>, such as X-ray imager <NUM> in <FIG> or first X-ray imager <NUM> or second X-ray imager <NUM> in <FIG>. Pixel detector element <NUM> includes a photodiode <NUM>, a reset switch <NUM>, a readout switch <NUM>, and a voltage follower <NUM>. As shown, photodiode <NUM> is communicatively coupled to a reference voltage Vref via reset switch <NUM> and is communicatively coupled to a data line <NUM> via readout switch <NUM> and voltage follower <NUM>.

Alternatively, in some embodiments, the positions of reference voltage Vref and a bias voltage <NUM> (e.g., ground) are reversed. In such embodiments, when reset switch <NUM> is closed, one side of photodiode <NUM> is connected to bias voltage <NUM>.

In operation, photodiode <NUM> produces a charge when photons that are generated by scintillator material <NUM> in <FIG> are incident on photodiode <NUM>. The charge accumulated by photodiode <NUM> is readout to data line <NUM> when readout switch <NUM> is closed and is converted to a digital signal by an analog-to-digital converter (ADC) <NUM> that is external to pixel detector element <NUM> and is coupled to data line <NUM>. If pixel detector element <NUM> includes voltage follower <NUM>, the charge accumulated by photodiode <NUM> is reset when reset switch <NUM> is closed.

<FIG> is a circuit diagram <NUM> of pixel detector element <NUM>, according to an example not forming part of the present invention. The embodiment of pixel detector element <NUM> illustrated in <FIG> is substantially similar to the embodiment of pixel detector element <NUM> illustrated in <FIG>, with the following exceptions. First, pixel detector element <NUM> does not include a voltage follower <NUM>. Second, photodiode <NUM> is not directly coupled to reference voltage Vref via reset switch <NUM>. Instead, photodiode <NUM> is coupled to reference voltage Vref via a reset switch <NUM> and data line <NUM> as shown. Alternatively, in some embodiments, reset switch <NUM> is integrated into ADC <NUM>. In either case, in such embodiments, readout switch <NUM> is closed whenever the reset switch <NUM> is closed so that the accumulation of signal in the photodiode during that time is prevented.

According to the embodiments illustrated in <FIG> and <FIG>, application of treatment beam pulses is restricted to treatment intervals and application of imaging beam pulses is restricted to imaging intervals, so that noise generated by the treatment beam pulses are prevented from being captured by photodiode <NUM>. Such embodiments are illustrated in <FIG> and <FIG>.

<FIG> is a schematic timing diagram <NUM> illustrating the application of treatment beam pulses <NUM> during treatment intervals <NUM> and the application of imaging beam pulses <NUM> during imaging intervals <NUM>, according to an embodiment of the present disclosure. Specifically, <FIG> applies to the embodiment of pixel detector element <NUM> illustrated in <FIG>. Also depicted in <FIG> is the timing of treatment beam pulses <NUM> and imaging beam pulses <NUM> relative to the opening and closing of reset switch <NUM>, readout switch <NUM>, and activation of ADC <NUM>.

During treatment intervals <NUM>, one or more treatment beam pulses <NUM> are directed to a target volume. In the embodiment illustrated in <FIG>, treatment interval <NUM> is shown with three treatment beam pulses <NUM>, but in practice, a treatment interval <NUM> can include up to one hundred or more treatment beam pulses. In some embodiments, a typical treatment beam pulse <NUM> (for example, of treatment beam <NUM> in <FIG>) is on the order of about <NUM> microseconds (µs) in duration, and is delivered about every <NUM>-<NUM> milliseconds (ms). In such embodiments, treatment beam pulses <NUM> are significantly shorter in duration than the imaging beam pulses <NUM> that occur during imaging intervals <NUM>. In addition, during treatment intervals <NUM>, reset switch <NUM> is closed, so that photodiode <NUM> is coupled to reference voltage Vref and cannot accrue charge. Thus, even though scattered X-rays from treatment beam pulses <NUM> strike scintillator layer <NUM> in <FIG>, which generates photons that are incident on photodiode <NUM>, photodiode <NUM> cannot accumulate such potential image noise as charge.

During each imaging interval <NUM>, one or more imaging beam pulses <NUM> are directed to the target volume. In the embodiment illustrated in <FIG>, a single imaging beam pulse <NUM> is illustrated. In some embodiments, a typical imaging beam pulse <NUM> (for example, of imaging X-rays <NUM> in <FIG>) is on the order of about <NUM> in duration, and is employed to generate a single projection image of the target volume. In such embodiments, one such image is acquired approximately every <NUM>-<NUM>. The one or more imaging beam pulses <NUM> occur during an irradiation portion <NUM> of each imaging interval <NUM>, and signal accumulated in each pixel detector element <NUM> is read out from each pixel in a readout portion <NUM> of each imaging interval <NUM>. As shown, reset switch <NUM> is opened during imaging intervals <NUM> so that photons generated by imaging beam pulses <NUM> and incident on photodiode <NUM> cause an image signal to accumulate in photodiode <NUM>. Further, readout switch <NUM> is closed during each imaging interval <NUM>, for example in synchronization with other photodiodes (not shown) of the X-ray imager and with activation of ADC <NUM>. In this way, an image signal is read out from each photodiode <NUM> of the X-ray imager during readout portion <NUM> of each imaging interval <NUM>.

As shown, treatment intervals <NUM> and imaging intervals <NUM> do not overlap in time, and instead are interleaved in the time domain. Thus, scattered X-rays that occur during treatment intervals <NUM> cannot be registered as charge by photodiode <NUM>. In the embodiment illustrated in <FIG>, a duration <NUM> of imaging intervals <NUM> is greater than a time interval <NUM> between two treatment beam pulses <NUM>. As a result, in embodiments in which treatment beam pulses <NUM> are directed through the target volume at regular intervals, one or more treatment beam pulses <NUM> are necessarily timed to occur during each imaging interval <NUM>. In such embodiments, such treatment beam pulses <NUM> are inhibited from being directed through the target volume so that photodiode <NUM> does not accumulate charge caused by scattered X-rays. Inhibited treatment beam pulses <NUM> are indicated with dashed lines. In the embodiment illustrated in <FIG>, imaging intervals <NUM> are shown to include a single inhibited treatment beam pulse <NUM>. In practice, a single imaging interval <NUM> can include anywhere from <NUM> inhibited treatment beam pulses <NUM> up to <NUM> or more, depending on the duration <NUM> of imaging intervals <NUM> and time interval <NUM> between two treatment beam pulses <NUM>.

In some embodiments, treatment control computer <NUM> (shown in <FIG>) performs the logical operations for determining whether the reset switches <NUM> for photodiodes <NUM> are open, indicating that an imaging interval <NUM> is underway. In such embodiments, treatment control computer <NUM> then prevents LINAC <NUM> from generating the inhibited treatment beam pulse or pulses <NUM>. Further, in some embodiments, dosing logic included in treatment control computer <NUM> can cause additional treatment beam pulses <NUM> to be applied to the target volume at the end of a treatment fraction to recover dose lost by the elimination of inhibited treatment beam pulses <NUM>.

In the embodiment illustrated in <FIG>, duration <NUM> of imaging intervals <NUM> is greater than time interval <NUM> between two treatment beam pulses <NUM>, and one or more treatment beam pulses <NUM> are timed to occur during each imaging interval <NUM>. By reducing duration <NUM> of imaging intervals <NUM>, the number of treatment beam pulses <NUM> that need to be inhibited can be reduced or eliminated, thereby preventing or reducing dose lost due to inhibited treatment beam pulses <NUM>. In some embodiments, a duration <NUM> of each irradiation portion <NUM> is reduced by delivering a higher-power imaging beam pulse <NUM> during the irradiation portion <NUM>. Thus, in such embodiments, the X-ray imager of RT system <NUM> includes an X-ray tube having a power of up to about <NUM> kW.

Alternatively or additionally, in some embodiments, a duration <NUM> of each readout portion <NUM> is reduced by performing the readout operation of photodiode <NUM> more quickly. For example, the inclusion of voltage follower <NUM> in each pixel detector element <NUM> enables very fast readout of photodiode <NUM>, since the circuit of pixel detector element <NUM> is not restricted by the resistor-capacitor (RC) time constant of the circuit.

<FIG> is a schematic timing diagram <NUM> illustrating the application of treatment beam pulses <NUM> during treatment intervals <NUM> and the application of imaging beam pulses <NUM> during imaging intervals <NUM>, according to another embodiment of the present disclosure. Specifically, <FIG> applies to the embodiment of pixel detector element <NUM> illustrated in <FIG>. Timing diagram <NUM> is substantially similar to timing diagram <NUM> in <FIG>, except that, in such a configuration, readout switch <NUM> is closed whenever reset switch <NUM> is closed so that the accumulation of signal in photodiode <NUM> during that time is prevented.

<FIG> sets forth a flowchart of a radiation therapy process, 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. Although the method is described in conjunction with the systems of <FIG>, persons skilled in the art will understand that any suitably configured radiation therapy system is within the scope of the present disclosure.

Prior to the method steps, setup volumetric (3D) image data for a digital volume <NUM> that includes a target volume <NUM> is acquired. For example, the patient is positioned on couch <NUM> of RT system <NUM> and the setup volumetric image data is generated by, for example, CBCT image acquisition. The setup volumetric image data includes image information for each voxel <NUM> in digital volume <NUM>. When produced by a CBCT process, the setup volumetric image data can include hundreds of distinct digital X-ray projection images of digital volume <NUM>. Auto-segmentation and deformable registration of the digital volume is then performed, followed by patient position adjustment. Auto-segmentation includes the delineation of target volumes and organs at risk within digital volume <NUM>. Deformable registration adjusts contours generated in an earlier planning phase for target volume <NUM> and any organs at risk. The deformable registration process compensates for changes in the shape and relative location of target volume <NUM> and organs at risk, for example due to stomach, colon, and bladder filling, tumor shrinkage, and other factors. The current position of the patient is also adjusted, when applicable, to precisely align target volume <NUM> with the now modified planning target volume. For example, the position of couch <NUM> can be automatically and/or manually adjusted to align target volume <NUM> with the modified planning target volume.

A method <NUM> begins at step <NUM>, when performance of one treatment arc of the current RT fraction is initiated. In step <NUM> a patient breath-hold begins. In some embodiments, the RT treatment included multiple fractions (i.e., multiple treatment arcs). Alternatively, in other embodiments, the RT treatment consists of a single treatment arc.

In step <NUM>, treatment control computer <NUM> causes gantry <NUM> to rotate continuously in a first direction through the treatment arc.

In step <NUM>, while gantry <NUM> rotates continuously in the first direction, treatment control computer <NUM> begins directing a series of pulses of treatment X-rays to target volume <NUM>, such as treatment beam pulses <NUM> or <NUM>.

In step <NUM>, treatment control computer <NUM> selects the next treatment beam pulse to be directed to target volume <NUM>.

In step <NUM>, treatment control computer <NUM> determines whether the next imaging interval <NUM> is timed to begin before the selected next treatment beam pulse. If yes, method <NUM> proceeds to step <NUM>; if no, method <NUM> proceeds to step <NUM>.

In step <NUM>, treatment control computer <NUM> determines whether the selected next treatment beam pulse is timed to occur during an imaging interval <NUM>. If yes, method <NUM> proceeds to step <NUM>; if not method <NUM> proceeds to step <NUM>.

In step <NUM>, treatment control computer <NUM> causes the next selected treatment beam pulse to be directed to target volume <NUM>.

In step <NUM>, treatment control computer <NUM> determines whether there are any remaining treatment intervals <NUM> or imaging intervals <NUM> in the treatment arc. If yes, method <NUM> proceeds back to step <NUM>, and the next treatment beam pulse is selected; if no, method <NUM> proceeds to step <NUM> and method <NUM> terminates.

Step <NUM> is performed in response to treatment control computer <NUM> determining that the next imaging interval <NUM> is timed to begin before the selected next treatment beam pulse. In step <NUM>, treatment control computer <NUM> causes the next imaging interval <NUM> to begin. Specifically, reset switch <NUM> is opened during imaging intervals <NUM>, so that each photodiode <NUM> in the X-ray imager can accrue charge, and readout switch <NUM> is closed, so that an image signal is read out from the photodiodes <NUM>. While step <NUM> is being performed, method <NUM> proceeds to step <NUM>.

Step <NUM> is performed in response to treatment control computer <NUM> determining that the selected next treatment beam pulse is timed to occur during an imaging interval <NUM>. In step <NUM>, treatment control computer <NUM> inhibits one or more treatment beam pulses <NUM> from being directed through the target volume during the current imaging interval <NUM>.

In optional step <NUM>, treatment control computer <NUM> updates the series of treatment beam pulses <NUM> to be directed to target volume <NUM>.

In some embodiments, an imaging interval fits between a series of multiple treatment beam pulses and, as a result, inhibition of treatment beam pulses is not performed. One such embodiment is illustrated in <FIG>.

<FIG> is a schematic timing diagram <NUM> illustrating the timing of imaging beam pulses <NUM> during imaging intervals <NUM> that have a shorter duration <NUM> than a time interval <NUM> between two treatment beam pulses <NUM>, according to an embodiment of the present disclosure. Also depicted in <FIG> is the timing of treatment beam pulses <NUM> during treatment intervals <NUM>, the opening and closing of reset switch <NUM> and readout switch <NUM>, and activation of ADC <NUM>. As shown, time interval <NUM> is longer in duration than imaging intervals <NUM>. Therefore, inhibition of treatment beam pulses <NUM> is not performed.

In some embodiments, the duration of imaging intervals <NUM> can be shortened to be less than time interval <NUM> due to the abbreviated readout portion <NUM> enabled by voltage follower <NUM>. Alternatively or additionally, the duration of imaging intervals <NUM> can be shortened to be less than time interval <NUM> due to the abbreviated irradiation portion <NUM> enabled by a sufficiently high-power imaging X-ray source. Alternatively or additionally, time interval <NUM> separating each treatment beam pulse <NUM> can be lengthened, for example up to about <NUM>. Thus, in one example embodiment, an X-ray image is acquired in about <NUM> to <NUM>, and treatment beam pulses <NUM> are directed to the target volume every <NUM> (or more). Consequently, each imaging interval <NUM> can be executed between each treatment beam pulse <NUM>.

In some cases not forming part of the present invention, treatment X-rays are delivered to a target volume at the same time that imaging X-rays are also delivered to the target volume for generating volumetric image data of the target volume. That is, during an imaging interval in which imaging X-rays are received by an X-ray imaging device, one or more pulses of treatment X-rays are also directed to the target volume. However, in such cases, an electronic shutter included in each pixel of the X-ray imaging device prevents image signal from being accumulated during portions of the imaging interval in which the pulses of treatment X-rays are directed to the target volume. One such example is described below in conjunction with <FIG>.

<FIG> is a circuit diagram <NUM> of pixel detector element <NUM> that includes an electronic shutter <NUM>. In this case, the pixel detector element <NUM> enables the X-ray imager to receive imaging X-rays during an imaging interval in which pulses of treatment X-rays are also directed to the target volume. Specifically, electronic shutter <NUM> prevents charge from accumulating in a signal integrator <NUM> that is associated with pixel detector element <NUM> when the treatment X-rays are directed to the target volume.

Pixel detector element <NUM> includes a photodiode <NUM> electronic shutter <NUM>, a readout/reset switch <NUM>, and signal integrator <NUM>. In some cases, pixel detector element <NUM> further includes a capacitor reset switch <NUM>. The pixel detector element <NUM> illustrated in <FIG> can be included in an X-ray imager of RT system <NUM>, such as X-ray imager <NUM> in <FIG> or first X-ray imager <NUM> or second X-ray imager <NUM> in <FIG>. As shown, photodiode <NUM> is communicatively coupled to a reference voltage Vref via electronic shutter <NUM> and is communicatively coupled to a data line <NUM> via signal integrator <NUM> and readout/reset switch <NUM>. In some cases, Vref can be any suitable reference voltage, for example ground or 0V. Photodiode <NUM> is a diode that does require the ability to store sufficient charge to produce an image signal. For example, in some cases the photodiode <NUM> is configured as a PN diode that does not include an intrinsic semiconductor region for storing charge. Alternatively, photodiode <NUM> is configured as a PIN diode, which can store significantly more charge than a PN diode. Pixel detector elements in conventional X-ray imagers are commonly configured as PIN diodes. In some cases, signal integrator <NUM> is a voltage integrator, a current integrator, or a charge integrator. In some embodiments, is formed from a single transistor 1004A and a single capacitor 1004B as shown. Alternatively, any other technically feasible charge-integrating device can be employed as signal integrator <NUM>.

In operation, photodiode <NUM> produces a charge or image signal when photons that are generated by scintillator material <NUM> in <FIG> are incident on photodiode <NUM>. The charge is accumulated in signal integrator <NUM>, which is included in pixel detector element <NUM> and outside photodiode <NUM>. The charge accumulated by photodiode <NUM> is then readout to data line <NUM> when readout/reset switch <NUM> is closed and is converted to a digital signal by an ADC <NUM> that is external to pixel detector element <NUM> and is coupled to data line <NUM>. In some cases, capacitor reset switch <NUM> is also closed to reset signal integrator <NUM>. According to some examples, electronic shutter <NUM> is closed during time intervals in which treatment beam pulses are directed to target volume <NUM>, so that signal integrator <NUM> does not receive additional image signal caused by scattered treatment beam pulses. Consequently, noise generated by the treatment beam pulses is prevented from being captured by photodiode <NUM>, as is illustrated in <FIG>.

<FIG> is a schematic timing diagram <NUM> illustrating the application of treatment beam pulses <NUM> and an imaging beam pulse <NUM> during an imaging interval <NUM> Also depicted in <FIG> is the timing of the opening and closing of electronic shutter <NUM>, readout/reset switch <NUM>, and activation of ADC <NUM> relative to treatment beam pulses <NUM> and imaging beam pulses <NUM>.

Imaging interval <NUM> includes an irradiation portion <NUM> and a readout portion <NUM>. During irradiation portion <NUM>, at least one imaging beam pulse <NUM> is directed to the target volume and is received by an X-ray imager of RT system <NUM>. In some cases, a typical imaging beam pulse <NUM> (for example, of imaging X-rays <NUM> in <FIG>) is on the order of about <NUM> in duration, and is employed to generate a single projection image of the target volume. In such cases, one such image is acquired approximately every <NUM>-<NUM>. During readout portion <NUM>, for each pixel detector element <NUM> included in the X-ray imager, image signal stored in the signal integrator <NUM> of that pixel detector element <NUM> is read out to ADC <NUM>. Specifically, readout/reset switch <NUM> is closed during readout portion <NUM>, for example in synchronization with other pixel detector elements <NUM> (not shown) of the X-ray imager and with activation of ADC <NUM>. In this way, an image signal is read out from each signal integrator <NUM> of the X-ray imager during readout portion <NUM> of each imaging interval <NUM>.

As shown, before, during, and after imaging interval <NUM>, one or more treatment beam pulses <NUM> are directed to a target volume. In some cases, a typical treatment beam pulse <NUM> (for example, of treatment beam <NUM> in <FIG>) is on the order of about <NUM> in duration and is delivered about every <NUM>-<NUM>. According to the example illustrated in <FIG>, electronic shutter <NUM> is closed when a treatment beam pulse <NUM> is directed to target volume <NUM>. Consequently, while electronic shutter <NUM> is closed, an output of photodiode <NUM> is coupled to ground and signal integrator <NUM> cannot receive additional image signal from photodiode <NUM>. Thus, even though scattered X-rays from treatment beam pulses <NUM> strike scintillator layer <NUM> in <FIG> and generate photons that are incident on photodiode <NUM>, signal integrator <NUM> cannot accumulate such potential image noise as charge.

<FIG> sets forth a flowchart of a radiation therapy process, 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. Although the method is described in conjunction with the systems of <FIG>, <FIG>, and <FIG>, persons skilled in the art will understand that any suitably configured radiation therapy system is within the scope of the present disclosure.

Prior to the method steps, similar setup procedures are performed as described above in conjunction with method <NUM>, including acquiring volumetric image data for a digital volume <NUM>, performing auto-segmentation and deformable registration, and positioning the patient prior to the RT treatment.

A method <NUM> begins at step <NUM>, when performance of one treatment arc of the current RT fraction is initiated. In step <NUM> a patient breath-hold begins. In some embodiments, the RT treatment included multiple fractions. Alternatively, in other embodiments, the RT treatment consists of a single treatment arc.

In step <NUM>, treatment control computer <NUM> determines whether the selected next treatment beam pulse is timed to occur during an imaging interval <NUM>. If yes, method <NUM> proceeds to step <NUM>; if no, method <NUM> proceeds to step <NUM>.

In step <NUM>, treatment control computer <NUM> causes electronic shutter <NUM> in each pixel detector element <NUM> to close prior to directing the selected next treatment beam pulse to target volume <NUM>. For example, in some embodiments, treatment control computer <NUM> causes the output from photodetector <NUM> to be communicatively coupled to ground. Thus, treatment control computer <NUM> prevents an image signal from accumulating in each signal integrator <NUM> of the X-ray imager.

In step <NUM>, treatment control computer <NUM> causes the selected next treatment beam pulse to be directed to target volume <NUM>.

In step <NUM>, treatment control computer <NUM> causes electronic shutter <NUM> in each pixel detector element <NUM> to open. For example, in some embodiments, treatment control computer <NUM> causes the output from photodetector <NUM> to be communicatively decoupled from ground. Thus, treatment control computer <NUM> enables an image signal to again accumulate in each signal integrator <NUM> of the X-ray imager.

In step <NUM>, treatment control computer <NUM> determines whether there are any remaining treatment beam pulses in the treatment arc. If yes, method <NUM> proceeds back to step <NUM>, and the next treatment beam pulse is selected; if no, method <NUM> proceeds to step <NUM> and method <NUM> terminates.

In step <NUM>, treatment control computer <NUM> directs the selected next treatment beam pulse to target volume <NUM> and method <NUM> proceeds to step <NUM>.

By way of illustration, embodiments of the disclosure have been described with respect to an RT system that includes a circular gantry and generates volumetric image data for a target volume. However, various embodiments as described herein can also be beneficially implemented in RT systems that have other configurations and/or perform two-dimensional imaging of a target volume. For example, various embodiments of the present disclosure can be advantageously applied to a radiation system configured to perform fluoroscopy, serial imaging, or triggered imaging. In such embodiments, an X-ray imager, a treatment beam generator, and/or an imaging X-ray source may not be coupled to a rotating gantry, or may be employed while the gantry is stationary.

Implementation of the above-described embodiments enables an X-ray imager to collect imaging data of a target volume during IGRT without image noise caused by treatment beam scattering. Advantageously, motion or deformation of the target volume relative to the radiation therapy system and/or changes in the surrounding organs at risk can be more reliably detected and compensated for. Consequently, more accurate adaption to the shape and relative position of organs at risk and the target volume is enabled, which facilitates better therapeutic results.

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.

Claim 1:
A radiation treatment system comprising:
an X-ray imager (<NUM>);
a treatment-delivering X-ray source (<NUM>) configured to direct treatment X-rays (<NUM>) to a target volume (<NUM>);
an imaging X-ray source (<NUM>) configured to direct imaging X-rays (<NUM>) through the target volume and toward the X-ray imager; and
a controller configured to:
direct a first pulse from a series of pulses (<NUM>) of treatment X-rays to the target volume
after directing the first pulse of treatment X-rays to the target volume and during a first imaging interval (<NUM>) in which no treatment X-rays are directed to the target volume, direct first imaging X-rays through the target volume; and
after directing the first imaging X-rays through the target volume, direct a second pulse from the series of pulses (<NUM>) of treatment X-rays to the target volume, a time interval (<NUM>) between two treatment beam pulses (<NUM>) separating each treatment beam pulse;
wherein the radiation treatment system is configured to shorten a duration of an imaging interval to be less than the time interval between two treatment beam pulses by means of a voltage follower (<NUM>) communicatively coupled to an output of a photodiode (<NUM>) in a pixel detector element (<NUM>) of the X-ray imager;
wherein said photodiode is communicatively coupled to a data line (<NUM>) via a readout switch (<NUM>) and said voltage follower,
wherein the voltage follower is included in each pixel detector element (<NUM>) so that a circuit of the pixel detector element is not restricted by a resistor-capacitor time constant of said circuit, the voltage follower enabling an abbreviated readout portion that shortens the duration of imaging intervals; and
wherein said controller is further configured to communicatively connect the voltage follower to the data line during the read-out portion (<NUM>) of the first imaging interval (<NUM>).