U.S. patent application number 17/590804 was filed with the patent office on 2022-05-19 for reduction of image lag in an x-ray detector panel.
This patent application is currently assigned to Varian Medical Systems International AG. The applicant listed for this patent is Varian Medical Systems International AG. Invention is credited to Simon KELLER, Daniel MORF.
Application Number | 20220159198 17/590804 |
Document ID | / |
Family ID | 1000006113954 |
Filed Date | 2022-05-19 |
United States Patent
Application |
20220159198 |
Kind Code |
A1 |
KELLER; Simon ; et
al. |
May 19, 2022 |
REDUCTION OF IMAGE LAG IN AN X-RAY DETECTOR PANEL
Abstract
A radiation therapy system is configured with fast readout of
X-ray images with significantly reduced image lag. 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. 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.
Inventors: |
KELLER; Simon; (Villigen,
CH) ; MORF; Daniel; (Buch am Irchel, CH) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Varian Medical Systems International AG |
Palo Alto |
CA |
US |
|
|
Assignee: |
Varian Medical Systems
International AG
Palo Alto
CA
|
Family ID: |
1000006113954 |
Appl. No.: |
17/590804 |
Filed: |
February 1, 2022 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
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16714501 |
Dec 13, 2019 |
11245860 |
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17590804 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
H04N 5/3741 20130101;
H04N 5/3742 20130101; A61B 6/4233 20130101; H04N 5/32 20130101;
G01T 1/247 20130101 |
International
Class: |
H04N 5/32 20060101
H04N005/32; A61B 6/00 20060101 A61B006/00; G01T 1/24 20060101
G01T001/24; H04N 5/374 20060101 H04N005/374 |
Claims
1. A method of acquiring successive X-ray images of a target volume
with an X-ray detector panel, the method comprising: determining
that a first reset interval associated with a first X-ray image
acquisition of the target volume is completed, wherein, during the
first reset interval, at least a portion of residual charge is
transferred from each pixel in a first array of pixels of the X-ray
detector panel and at least a portion of residual charge is
transferred from each pixel in a second array of pixels of the
X-ray detector panel; after determining that the first reset
interval is completed, transferring first accumulated charge from
each pixel in the first array of pixels of the X-ray detector panel
to a readout device of the X-ray detector panel during a first
readout interval, wherein residual charge remains in each pixel in
the first array of pixels after transferring the first accumulated
charge; after the first readout interval, transferring second
accumulated charge from each pixel in the second array of pixels of
the X-ray detector panel to the readout device during a second
readout interval, wherein residual charge remains in each pixel in
the second array of pixels after transferring the second
accumulated charge; and during a second reset interval that follows
the second readout interval and is associated with a second X-ray
image acquisition of the target volume, concurrently transferring
at least a portion of the residual charge from each pixel in the
first array of pixels and at least a portion of the residual charge
from each pixel in the second array of pixels.
2. The method of claim 1, wherein, during the first reset interval:
the portion of residual charge transferred from each pixel in the
first array of pixels is transferred to a readout device of the
X-ray detector panel; and the portion of residual charge
transferred from each pixel in the second array of pixels is
transferred to the readout device of the X-ray detector panel.
3. The method of claim 1, wherein the first accumulated charge and
the second accumulated charge are associated with the second X-ray
image acquisition of the target volume.
4. The method of claim 1, wherein the first X-ray image acquisition
and the second X-ray image acquisition are included in a cone-beam
computed tomography process.
5. The method of claim 1, wherein the first X-ray image acquisition
and the second X-ray image acquisition are included in a digital
tomosynthesis process.
6. The method of claim 1, wherein the first X-ray image acquisition
comprises acquiring a first X-ray projection image of the target
volume and the second X-ray image acquisition comprises acquiring a
second X-ray projection image of the target volume.
7. The method of claim 1, further comprising, during the first
reset interval, concurrently transferring at least a portion of the
residual charge from each pixel in each array of pixels of the
X-ray detector panel to the reference voltage associated with the
readout device.
8. The method of claim 1, wherein concurrently transferring at
least the portion of the residual charge from each pixel in the
first array of pixels and at least the portion of the residual
charge from each pixel in the second array of pixels comprises
electrically coupling each pixel in the second array of pixels to a
reference voltage
9. The method of claim 8, wherein the reference voltage is
associated with the readout device.
10. The method of claim 1, wherein the first array of pixels
comprises a first row of pixels and the second array of pixels
comprises a second row of pixels.
11. The method of claim 1, wherein a start time of the first reset
interval is selected to occur prior to a treatment beam pulse and
an end time of the first reset interval is selected to occur after
the treatment beam pulse.
12. The method of claim 7, wherein a megavolt treatment beam is
directed to the target volume during the treatment beam pulse.
13. The method of claim 1, further comprising: prior to the first
readout interval, preparing each pixel in the first array of pixels
and each pixel in the second array of pixels to accumulate charge
in response to incident X-rays; and causing imaging X-rays to be
directed to the X-ray detector panel.
14. An apparatus comprising: an imaging X-ray source configured to
direct imaging X-rays through a target volume and toward an X-ray
detector panel; and a controller configured to: determine that a
first reset interval associated with a first X-ray image
acquisition of the target volume is completed, wherein, during the
first reset interval, at least a portion of residual charge is
transferred from each pixel in a first array of pixels of the X-ray
detector panel and at least a portion of residual charge is
transferred from each pixel in a second array of pixels of the
X-ray detector panel; after determining that the first reset
interval is completed, transfer first accumulated charge from each
pixel in the first array of pixels of the X-ray detector panel to a
readout device of the X-ray detector panel during a first readout
interval, wherein residual charge remains in each pixel in the
first array of pixels after transferring the first accumulated
charge; after the first readout interval, transfer second
accumulated charge from each pixel in the second array of pixels of
the X-ray detector panel to the readout device during a second
readout interval, wherein residual charge remains in each pixel in
the second array of pixels after the second accumulated charge is
transferred; and during a second reset interval that follows the
second readout interval and is associated with a second X-ray image
acquisition of the target volume, concurrently transfer at least a
portion of the residual charge from each pixel in the first array
of pixels and at least a portion of the residual charge from each
pixel in the second array of pixels.
15. The apparatus of claim 1, wherein the first array of pixels
comprises a first row of pixels of the X-ray detector panel and the
second array of pixels comprises a second row of pixels of the
X-ray detector panel.
16. The apparatus of claim 14, wherein a start time of the second
reset interval is selected to occur prior to a treatment beam pulse
and an end time of the second reset interval is selected to occur
after the treatment beam pulse.
17. The apparatus of claim 16, further comprising linear
accelerator configured to generate a megavolt treatment beam,
wherein the megavolt treatment beam is directed to the target
volume during the treatment beam pulse.
18. The apparatus of claim 14, wherein the first accumulated charge
and the second accumulated charge are associated with the second
X-ray image acquisition of the target volume.
19. The apparatus of claim 14, wherein the first X-ray image
acquisition and the second X-ray image acquisition are included in
a cone-beam computed tomography process.
20. The apparatus of claim 14, wherein the first X-ray image
acquisition and the second X-ray image acquisition are included in
a digital tomosynthesis process.
Description
CROSS-REFERENCE TO RELATED APPLICATION
[0001] The present application is a continuation under 35 U.S.C.
.sctn. 120 of U.S. patent application Ser. No. 16/714,501, filed
Dec. 13, 2019. The aforementioned U.S. Patent Application,
including any appendices or attachments thereof, is hereby
incorporated by reference in its entirety.
BACKGROUND
[0002] Unless otherwise indicated herein, the approaches described
in this section are not prior art to the claims in this application
and are not admitted to be prior art by inclusion in this
section.
[0003] 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.
[0004] 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.
SUMMARY
[0005] 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.
[0006] The foregoing summary is illustrative only and is not
intended to be in any way limiting. In addition to the illustrative
aspects, embodiments, and features described above, further
aspects, embodiments, and features will become apparent by
reference to the drawings and the following detailed
description.
BRIEF DESCRIPTION OF THE DRAWINGS
[0007] The foregoing 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.
[0008] FIG. 1 is a perspective view of a radiation therapy system
that can beneficially implement various aspects of the present
disclosure.
[0009] FIG. 2 schematically illustrates a drive stand and a gantry
of the radiation system of FIG. 1, according to various embodiments
of the current disclosure.
[0010] FIG. 3 schematically illustrates a cross-sectional view of
an X-ray imager, according to one embodiment of the disclosure.
[0011] FIG. 4 schematically illustrates a digital volume that is
constructed based on projection images generated by one or more
X-ray imagers included in the RT system of FIG. 1, according to
various embodiments of the current disclosure.
[0012] FIG. 5 is a partial circuit diagram of a photosensitive
element and detector circuitry layer included in the X-ray imager
of FIG. 3, according to one embodiment of the disclosure.
[0013] FIG. 6 is a partial circuit diagram of the photosensitive
element and detector circuitry layer included in the X-ray imager
of FIG. 3, according to another embodiment of the disclosure.
[0014] FIG. 7 is a partial circuit diagram of the photosensitive
element and detector circuitry layer included in the X-ray imager
of FIG. 3, according to yet another embodiment of the
disclosure.
[0015] FIG. 8 is a timing diagram schematically illustrating charge
accumulation and loss in pixel detector elements during an
irradiation phase, a readout phase, and a reset phase of a single
X-ray image acquisition, according to an embodiment of the present
disclosure.
[0016] FIG. 9 is a schematic timing diagram illustrating charge
loss for two adjacent rows of pixel detector elements during the
readout phase of FIG. 8, according to an embodiment of the present
disclosure.
[0017] FIG. 10 is a schematic timing diagram illustrating charge
loss for two adjacent rows of pixel detector elements during the
reset phase of FIG. 8, according to an embodiment of the present
disclosure.
[0018] FIG. 11 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.
[0019] FIG. 12A is a timing diagram schematically illustrating
charge loss for different rows of pixel detector elements during a
reset phase, according to an embodiment of the present
disclosure.
[0020] FIG. 12B illustrates a portion of the reset phase of FIG.
12A for a first array of pixel detector elements and an adjacent
second array of pixel detector elements, according to an embodiment
of the present disclosure.
[0021] FIG. 13 is a timing diagram schematically illustrating the
relative timing of an irradiation phase, a readout phase, and a
reset phase relative to the delivery of a treatment beam pulse
1350, according to an embodiment of the present disclosure.
DETAILED DESCRIPTION
[0022] In the following detailed description, reference is made to
the accompanying drawings, which form a part 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 spirit or 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.
INTRODUCTION
[0023] 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.
[0024] 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
laparoscopic 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.
[0025] 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.
[0026] 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.
[0027] 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.
[0028] 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.
1.
System Overview
[0029] FIG. 1 is a perspective view of a radiation therapy system
100 that can beneficially implement various aspects of the present
disclosure. Radiation therapy (RT) system 100 is a radiation system
configured to detect intra-fraction motion in near-real time using
X-ray imaging techniques. Thus, RT system 100 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 100 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 100 is described herein configured with a
circular gantry. In other embodiments, radiation therapy system 100
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.
[0030] Generally, RT system 100 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 100 may include one or more touchscreens 101, couch
motion controls 102, a bore 103, a base positioning assembly 105, a
couch 107 disposed on base positioning assembly 105, and an image
acquisition and treatment control computer 106, all of which are
disposed within a treatment room. RT system 100 further includes a
remote control console 110, which is disposed outside the treatment
room and enables treatment delivery and patient monitoring from a
remote location. Base positioning assembly 105 is configured to
precisely position couch 107 with respect to bore 103, and motion
controls 102 include input devices, such as button and/or switches,
that enable a user to operate base positioning assembly 105 to
automatically and precisely position couch 107 to a predetermined
location with respect to bore 103. Motion controls 102 also enable
a user to manually position couch 107 to a predetermined location.
In some embodiments, RT system 100 further includes one or more
cameras (not shown) in the treatment room for patient
monitoring.
[0031] FIG. 2 schematically illustrates a drive stand 200 and
gantry 210 of RT system 100, according to various embodiments of
the current disclosure. Covers, base positioning assembly 105,
couch 107, and other components of RT system 100 are omitted in
FIG. 2 for clarity. Drive stand 200 is a fixed support structure
for components of RT treatment system 110, including gantry 210 and
a drive system 201 for rotatably moving gantry 210. Drive stand 200
rests on and/or is fixed to a support surface that is external to
RT treatment system 110, such as a floor of an RT treatment
facility. Gantry 210 is rotationally coupled to drive stand 200 and
is a support structure on which various components of RT system 100
are mounted, including a linear accelerator (LINAC) 204, an MV
electronic portal imaging device (EPID) 205, an imaging X-ray
source 206, and an X-ray imager 207. During operation of RT
treatment system 110, gantry 210 rotates about bore 103 when
actuated by drive system 201.
[0032] Drive system 201 rotationally actuates gantry 210. In some
embodiments, drive system 201 includes a linear motor that can be
fixed to drive stand 200 and interacts with a magnetic track (not
shown) mounted on gantry 210. In other embodiments, drive system
201 includes another suitable drive mechanism for precisely
rotating gantry 210 about bore 201. LINAC 204 generates an MV
treatment beam 230 of high energy X-rays (or in some embodiments
electrons) and EPID 205 is configured to acquire X-ray images with
treatment beam 230. Imaging X-ray source 206 is configured to
direct a conical beam of X-rays, referred to herein as imaging
X-rays 231, through an isocenter 203 of RT system 100 to X-ray
imager 207, and isocenter 203 typically corresponds to the location
of a target volume 209 to be treated. In the embodiment illustrated
in FIG. 2, X-ray imager 207 is depicted as a planar device, whereas
in other embodiments, X-ray imager 207 can have a curved
configuration.
[0033] X-ray imager 207 receives imaging X-rays 231 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 209. 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 207. CBCT is typically employed to acquire projection
images over a relatively long acquisition arc, for example over a
rotation of 180.degree. or more of gantry 210. 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
230 to generate a 3D reconstruction confirming that target volume
209 has not moved or changed shape.
[0034] In some embodiments, partial-data reconstruction may be
performed by RT system 100 during portions of an IGRT process in
which partial image data is employed to generate a 3D
reconstruction of target volume 209. For example, as treatment beam
230 is directed to isocenter 203 while gantry 210 rotates through a
treatment arc, DTS image acquisitions can be performed to generate
image data for target volume 209. Because DTS image acquisition is
performed over a relatively short acquisition arc, for example
between about 10.degree. and 60.degree., near real-time feedback
for the shape and position of target volume 209 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 209. 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 209, either
for CBCT or DTS image acquisition.
[0035] In some embodiments, X-ray imager 207 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 207 is
configured as an indirect flat panel detector, a scintillator
material in X-ray imager 207 is excited by incident X-rays and
emits light, which is detected by a plurality of photodiodes. Each
photodiode generates a signal (e.g., an accumulated voltage 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 207 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 207 is illustrated in FIG. 3.
[0036] FIG. 3 schematically illustrates a cross-sectional view of
X-ray imager 207, according to one embodiment of the disclosure. As
shown, X-ray imager 207 includes a photosensitive element and
detector circuitry layer 301 formed on a substrate 302. In
addition, X-ray imager 207 includes a layer of scintillator
material 303 formed on photosensitive element and detector
circuitry layer 301. Also shown are incident X-rays 309 that have
passed through a patient, sample, or other object of interest after
being generated by imaging X-ray source 206 (shown in FIG. 2).
Together, photosensitive element and detector circuitry layer 301,
substrate 302, and scintillator material 303 form an X-ray imaging
array 305. It is noted that photosensitive element and detector
circuitry layer 301 is generally formed from a plurality of
processing layers, and that X-ray imaging array 305 may include
additional material layers not illustrated in FIG. 3.
[0037] Photosensitive element and detector circuitry layer 301
generally includes a plurality of pixel detector elements 310. Each
pixel detector element 310 includes a photosensitive element, such
as a photodiode, a photogate, or a phototransistor, as well as any
other circuitry suitable for operation as a pixel detector element
in X-ray imager 207. In some embodiments, the photosensitive
elements of pixel detector element 310 are amorphous silicon-based
semiconductor devices. Photosensitive element and detector
circuitry layer 301 may also include thin-film transistors (TFTs)
for reading out the digital signals from pixel detector elements
310. Scintillator material 303 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 (CsI), or cesium iodide thallium
(CsI:Tl)), among others.
[0038] In the embodiment illustrated in FIG. 3, X-ray imager 207 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 207 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.
[0039] In the embodiment illustrated in FIG. 2, RT system 100
includes a single X-ray imager and a single corresponding imaging
X-ray source. In other embodiments, RT system 100 can include two
or more X-ray imagers, each with a corresponding imaging X-ray
source. Thus, in such embodiments, RT system 100 includes a first
imaging X-ray source and a corresponding X-ray imager mounted on
gantry 210 and a second imaging X-ray source and corresponding
X-ray imager mounted on gantry 210. In such embodiments, the
inclusion of multiple X-ray imagers in RT system 100 facilitates
the generation of projection images (for reconstructing the target
volume) over a shorter image acquisition arc. For instance, when RT
system 100 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 100
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 100 includes a single imaging X-ray source, the
single imaging X-ray source can be configured as a multi-energy
source.
[0040] The projection images generated by X-ray imager 207 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. 4.
[0041] FIG. 4 schematically illustrates a digital volume 400 that
is constructed based on projection images generated by one or more
X-ray imagers included in RT system 100, 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 207, and in other embodiments
the projection images can be generated by multiple X-ray
imagers.
[0042] Digital volume 400 includes a plurality of voxels 401
(dashed lines) of anatomical image data, where each voxel 401
corresponds to a different location within digital volume 400. For
clarity, only a single voxel 401 is shown in FIG. 4. Digital volume
400 corresponds to a 3D region that includes target volume 410. In
FIG. 4, digital volume 400 is depicted as an 8.times.8.times.8
voxel cube, but in practice, digital volume 400 generally includes
many more voxels, for example orders of magnitude more than are
shown in FIG. 4.
[0043] For purposes of discussion, target volume 410 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.
[0044] According to various embodiments described below, image
information associated with each voxel 401 of digital volume 400 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 401 of digital volume 400 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 410 can be confirmed while the treatment is underway.
Thus, if a sufficient portion of the target volume 410 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.
[0045] FIG. 5 is a partial circuit diagram 500 of photosensitive
element and detector circuitry layer 301, according to one
embodiment of the disclosure. Photosensitive element and detector
circuitry layer 301 can be included in a suitable X-ray detector
panel, such as X-ray imager 207 in FIG. 2. As shown, photosensitive
element and detector circuitry layer 301 includes a plurality of
pixel detector elements 310 that are each communicatively coupled
to a readout stage 520 via a common data line 510. Generally,
photosensitive element and detector circuitry layer 301 includes an
M.times.N matrix of pixel detector elements 310, where M equals the
number of rows of pixel detector elements 310 and N equals the
number of columns of pixel detector elements 310. Generally, M and
N have values on the order of about 1000 to 4000 in an X-ray
detector panel. For clarity, only a single array of pixel detector
elements 310 that forms one column of M pixel detector elements 310
is shown in FIG. 5. In practice, photosensitive element and
detector circuitry layer 301 includes a total of N such arrays of
pixel detector elements 310, where each array forms one of the M
rows of pixels of an X-ray detector panel.
[0046] In the embodiment illustrated in FIG. 5, each pixel detector
element 310 includes a photodiode 501 and a readout switch 502 that
communicatively couples the photodiode 501 to readout stage 520. As
shown, each photodiode 501 is communicatively coupled to a bias
voltage V.sub.bias and is also communicatively coupled to data line
510 via a respective readout switch 502. Each readout switch 502 is
typically formed as part of the associated pixel detector element
310. Alternatively, in some embodiments, each readout switch 502 is
formed proximate to the associated pixel detector element 310. In
either case, readout switches 502 are generally formed as part of
photosensitive element and detector circuitry layer 301. For
example, in some embodiments, readout switches 502 are implemented
as thin-film transistors (TFTs) that are formed on the same
substrate as the pixel detector elements 310.
[0047] Readout stage 520 is a readout device configured to readout
accumulated charge from the pixel detector elements 310 of a
particular column 550 of pixel detector elements 310. Readout stage
520 reads out a particular pixel detector element 310 (for example,
the pixel detector element 310 associated with Pixel M-2) when the
readout switch 502 for that particular pixel detector element 310
(for example, readout switch 502 associated with Pixel M-2) closes
and communicatively couples that pixel detector element 310 to data
line 510. In operation, for each other column (not shown) of
photosensitive element and detector circuitry layer 301, a single
pixel detector element 310 can be simultaneously readout by the
readout stage 520 associated with the column. Thus, a complete row
of pixel detector elements 310 of photosensitive element and
detector circuitry layer 301 can be read out at one time by the
appropriately timed closing of one readout switch 502 in each
column of pixel detector elements 310.
[0048] Readout stage 520 is configured to convert analog signals,
such as charge accumulated in pixel detector elements 310, to
digital X-ray image signals. In some embodiments, readout stage 520
includes conversion circuitry 521 for converting such signals to
digital X-ray image signals. Conversion circuitry 521 can include
any technically feasible circuitry suitable for performing such
conversions. For example, in some embodiments, conversion circuitry
521 includes an analog-to-digital converter, an analog front-end,
or the like. In addition, in the embodiment illustrated in FIG. 5,
readout stage 520 includes a reset switch 522 that is configured to
communicatively couple data line 510 to a reference voltage
V.sub.ref, such as ground or any other suitable reference voltage.
Thus, when a readout switch 502 of a particular pixel detector
element 310 is closed while reset switch 522 is closed, charge
currently accumulated in that particular pixel detector element 310
discharges down to the reference voltage V.sub.ref. By contrast,
when a readout switch 502 of a particular pixel detector element
310 is closed while reset switch 522 is open, charge currently
accumulated in that particular pixel detector element 310
discharges to the readout stage 520 and is read by conversion
circuitry 521.
[0049] In the embodiment illustrated in FIG. 5, reset switch 522 is
included in readout stage 520. In other embodiments, a reset switch
is still communicatively coupled to data line 510, but is
implemented outside of readout stage 520. One such embodiment is
illustrated in FIG. 6. FIG. 6 is a partial circuit diagram 600 of
photosensitive element and detector circuitry layer 301, according
to another embodiment of the disclosure. As shown, in the
embodiment illustrated in FIG. 6, a reset switch 622 is configured
to selectively couple data line 510 to reference voltage V.sub.ref,
but is located at some other position along data line 510 than
readout stage 520. In yet other embodiments, a reset switch is
communicatively coupled to data line 510, but is implemented via
existing circuitry included in conversion circuitry 521. One such
embodiment is illustrated in FIG. 7. FIG. 7 is a partial circuit
diagram 700 of photosensitive element and detector circuitry layer
301, according to yet another embodiment of the disclosure. As
shown, in the embodiment illustrated in FIG. 7, a reset switch 722
is configured to selectively couple data line 510 to reference
voltage V.sub.ref and is included within conversion circuitry 521.
Alternatively, the functionality of reset switch 722 is implemented
by one or more components of conversion circuitry 521, and reset
switch 722 is not a dedicated switch or transistor for coupling
data line 510 to reference voltage V.sub.ref.
Reduction of Image Lag Via Pixel Charge Reset
[0050] According to various embodiments described herein, the
acquisition of an X-ray image with an X-ray imager, such as X-ray
imager 207, is performed in three phases: an irradiation phase, a
readout phase, and a reset phase. In the irradiation phase, each
pixel detector element 310 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 310 is transferred to readout stage 520 and
is processed. In the reset phase, residual charge is transferred
from each pixel detector element 310, thereby minimizing or
otherwise reducing image lag present in the next X-ray image to be
acquired. One such embodiment is described below in conjunction
with FIG. 8.
[0051] FIG. 8 is a timing diagram 800 schematically illustrating
charge accumulation and loss in pixel detector elements 310 during
an irradiation phase 810, a readout phase 820, and a reset phase
830 of a single X-ray image acquisition, according to an embodiment
of the present disclosure. More specifically, charge accumulation
and loss is shown for M rows of pixel detector elements 310 in an
X-ray detector panel that includes an M.times.N matrix of pixel
detector elements 310. One such row is described above in
conjunction with FIG. 5.
[0052] As shown, throughout irradiation phase 810, each of the M
rows are simultaneously irradiated, and charge is accumulated in
the pixel detector elements 310 of each row. In readout phase 820,
charge accumulated in the pixel detector elements 310 of each row
are read out sequentially. That is, the pixel detector elements 310
for row 1 are read out by readout stage 520, then the pixel
detector elements 310 for row 2 are read out by readout stage 520,
and so on, until the pixel detector elements 310 of all M rows are
read out and an X-ray image can be generated. In reset phase 830,
residual charge that remains in the pixel detector elements 310 of
each row is concurrently transferred from all rows of an X-ray
detector panel. Readout phase 820 is described in greater detail
below in conjunction with FIG. 9, and reset phase 830 is described
in greater detail below in conjunction with FIG. 10.
[0053] FIG. 9 is a schematic timing diagram 900 illustrating charge
loss for two adjacent rows of pixel detector elements 310 during
readout phase 820, according to an embodiment of the present
disclosure. In FIG. 9, a portion of readout phase 820 is shown for
a first array (row M-4) of pixel detector elements 310 and an
adjacent second array (row M-3) of pixel detector elements 310.
During readout phase 820, accumulated charge in each of the pixel
detector elements 310 of row M-4 is read during a first readout
time interval t.sub.read1, and accumulated charge in each of the
pixel detector elements 310 of row M-3 is read during a second
readout time interval t.sub.read2, that follows first readout time
interval t.sub.read1.
[0054] First readout time interval t.sub.read1 begins after the
accumulated charge in each of the pixel detector elements 310 of
the preceding row (e.g., row M-5, not shown) has been read out,
since photosensitive element and detector circuitry layer 301 is
typically connected to a readout stage 520 that is configured to
read out one row of pixel detector elements 310 at a time.
Similarly, second readout time interval t.sub.read2 begins after
the accumulated charge in each of the pixel detector elements 310
of row M-4 has been read out and first readout time interval tread1
has ended.
[0055] During first readout time interval t.sub.read1, the readout
switch 502 for each pixel detector element 310 of row M-4 closes
(indicated by ON state in FIG. 9) and an accumulated charge
Q.sub.M-4 decreases in magnitude over first readout time interval
t.sub.read1 from an initial charge value Q.sup.0.sub.M-4 to a
remainder charge value Q.sup.Rem.sub.M-4. The readout switch 502
for each pixel detector element 310 of row M-4 then opens
(indicated by OFF state in FIG. 9) and no more charge is read out
from the pixel detector element 310 of row M-4. Typically, the
value of remainder charge value Q.sup.Rem.sub.M-4 is a function of
the duration of first readout time interval t.sub.read1, a time
dependent release of trapped charges, and a pixel time constant
T.sub.pix, which is substantially the same for each photodiode 501
of an X-ray imager. For example, in some embodiments, the value of
remainder charge value Q.sup.Rem.sub.M-4 can be determined based on
Equation 1:
Q M - 4 Rem = Q M - 4 O .times. exp .function. ( - t read .times.
.times. 1 / T pix ) + Q M - 4 trap .function. ( t ) ( 1 )
##EQU00001##
[0056] Similarly, during second readout time interval t.sub.read2,
the readout switch 502 for each pixel detector element 310 of row
M-3 closes (indicated by ON state in FIG. 9) and an accumulated
charge Q.sub.M-3 decreases in magnitude over second readout time
interval t.sub.read2 from an initial charge value Q0M-3 to a
remainder charge value Q.sup.Rem.sub.M-3. The readout switch 502
for each pixel detector element 310 of row M-3 then opens
(indicated by OFF state in FIG. 9) and no more charge is read out
from the pixel detector element 310 of row M-3. This process
continues sequentially through the remaining rows of the X-ray
imager.
[0057] Charge cannot escape from a photodiode 501 once the
associated readout switch 502 opens. Therefore, in a conventional
X-ray imager, the value of remainder charge value Q.sup.Rem for a
row corresponds to image lag for each of the pixel detector element
310 of the row, since such remaining charge is present when the
subsequent irradiation phase 810 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
Q.sup.Rem 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).times.(number of rows of pixel detector elements).
Further, because the rate at which remainder charge value Q.sup.Rem
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 Q.sup.Rem. By
contrast, according to various embodiments described herein, reset
phase 830 enables remainder charge value Q.sup.Rem 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 810.
Consequently, image lag can be prevented without slowing panel
readout time by more than the duration of reset phase 830.
[0058] FIG. 10 is a schematic timing diagram 1000 illustrating
charge loss for two adjacent rows of pixel detector elements 310
during reset phase 830, according to an embodiment of the present
disclosure. In FIG. 10, a portion of reset phase 830 is shown for a
first array (row M-4) of pixel detector elements 310 and an
adjacent second array (row M-3) of pixel detector elements 310.
Reset phase 830 begins after readout phase 820 for all M rows of
the X-ray detector panel has been completed.
[0059] During reset phase 830, residual charge in each of the pixel
detector elements 310 is concurrently transferred from each of the
M rows of the X-ray detector panel, for example during a single
reset time interval t.sub.reset. Thus, during reset time interval
t.sub.reset, residual charge in each of the pixel detector elements
310 of row M-4 is transferred from the pixel detector elements, for
example to ground or some other suitable reference voltage
V.sub.ref. In addition, during reset time interval t.sub.reset,
residual charge in each of the pixel detector elements 310 of row
M-3 is transferred from the pixel detector elements to reference
voltage V.sub.ref. Further, during reset time interval t.sub.reset,
residual charge in each of the pixel detector elements 310 of the
remaining rows (not shown) of the X-ray detector panel is
transferred from the pixel detector elements to reference voltage
V.sub.ret. Thus, during reset time interval t.sub.reset, residual
charge from some or all of the M rows is concurrently transferred
from pixel detector elements 310 to reference voltage V.sub.ref. As
a result, the magnitude of image lag associated with each pixel
detector element 310 can be reduced much more quickly than by
increasing the duration of the readout time interval t.sub.read for
each row of pixel detector elements 310.
[0060] As shown in FIG. 10, at the beginning of reset time interval
t.sub.reset, residual charge in each particular pixel detector
element 310 of row M-4 equals the remainder charge value
Q.sup.Rem.sub.M-4, and at the end of reset time interval
t.sub.reset, residual charge in each particular pixel detector
element 310 of row M-4 equals a final charge value
Q.sup.Fin.sub.M-4 for that particular pixel detector element 310.
Similarly, at the beginning of reset time interval t.sub.reset,
residual charge in each particular pixel detector element 310 of
row M-3 equals the remainder charge value Q.sup.Rem.sub.M-3 of that
particular pixel detector element 310, and at the end of reset time
interval t.sub.reset, residual charge in each particular pixel
detector element 310 of row M-3 equals a final charge value
Q.sup.Fin.sub.M-3 for that particular pixel detector element 310.
Similar to remainder charge value Q.sup.Rem, in some embodiments, a
final charge value Q.sup.Fin for each pixel detector element 310 of
a particular row (for example row M-3) can be determined based on
Equation 2:
Q M - 3 Fin = QO M - 3 .times. exp .function. ( - t reset / Tpix )
+ Q M - 4 trap .function. ( t ) ( 2 ) ##EQU00002##
[0061] Thus, the magnitude of Q.sup.Fin for pixel detector elements
310 decays at substantially the same rate during reset time
interval t.sub.reset reset phase 830 as the magnitude of Q.sup.Rem
decays of during one of the readout time intervals of readout phase
820, such as readout time interval t.sub.read1. However, in reset
phase 830, charge is transferred from some, most, or all of the M
rows of the pixel detector elements 310 of the X-ray detector panel
simultaneously. Thus, increasing the duration of reset time
interval t.sub.reset can greatly reduce the magnitude of
accumulated charge (from Q.sup.Rem to Q.sup.Fin) in each of the
pixel detector elements 310 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 Q.sup.Fin) by increasing the duration
of each of the readout time intervals of readout phase 820, the
panel readout time for the X-ray detector panel is greatly
increased. For example, in one embodiment, reset time interval
t.sub.reset of reset phase 830 is set to be three times as long as
the readout time intervals of readout phase 820. Therefore, panel
readout time is increased by 3.times.(readout time interval
t.sub.reset). To achieve the same reduction in accumulated charge
in the pixel detector elements 310 of the X-ray detector panel
without using the reset phase 830, the panel readout time is
quadrupled, which significantly slows imaging framerate of the
X-ray detector.
[0062] FIG. 11 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 1111-1135. 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 FIGS. 1-10, persons skilled in the art will
understand that any suitably configured radiation therapy system is
within the scope of the present disclosure. 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 106 of RT system 100 (shown in
FIG. 1), a controller included in X-ray imager 207 (shown in FIG.
2), any other suitable controller associated with RT system 100, or
any combination thereof.
[0063] A method 1100 begins at step 1111, where treatment control
computer 106, a controller included in X-ray imager 207, or any
other suitable controller associated with RT system 100 causes
X-ray imager 207 to be prepared for irradiation phase 810 and the
acquisition of an X-ray image. In some embodiments, the controller
causes charge currently accumulated in pixel detector elements 310
of X-ray imager 207 to be transferred to ground or other reference
voltage V.sub.ref. For example, the controller causes readout
switch 502 of each pixel detector element 310 and reset switch 522
to close for a certain time interval to reduce accumulated charge
in pixel detector elements 310. The controller then causes readout
switch 502 of each pixel detector element 310 and reset switch 522
to open.
[0064] In step 1112, the controller begins irradiation phase 810.
For example, in some embodiments, the controller causes imaging
X-ray source 206 to direct imaging X-rays 231 through isocenter 203
of RT system 100 to X-ray imager 207.
[0065] In step 1113, X-ray imager 207 receives imaging X-rays 231
and charge is accumulated in some or all of pixel detector elements
310.
[0066] In step 1114, the controller ends irradiation phase 810. For
example, in some embodiments, the controller causes imaging X-ray
source 206 to stop directing imaging X-rays 231 to X-ray imager
207.
[0067] In step 1121, the controller selects a pixel array from the
M pixel arrays of X-ray imager 207 for readout. In some
embodiments, each of the M pixel arrays is configured as a row of
pixel detector elements 310. In other embodiments, each of the M
pixel arrays is configured as a column of pixel detector elements
310. In yet other embodiments, each of the M pixel arrays is
configured as any other group of pixel detector elements 310 that
are simultaneously read out by readout stage 520, such as a group
of pixel detector elements 310 that are located in a particular
region of X-ray imager 207. For clarity, method 1100 is described
herein in terms of rows of pixel detector elements 310 included in
X-ray imager 207, but method 1100 is equally applicable to any
other suitable pixel array configuration of pixel detector elements
310, such as columns or other groups of pixel detector elements
310.
[0068] In step 1122, the controller causes accumulated charge from
the pixel detector elements 310 of the selected pixel array to be
read out during a readout interval. For example, in some
embodiments, the controller causes the readout switch 502 of each
pixel detector element 310 in the selected pixel array to close for
a readout time interval t.sub.read while reset switch 522 to
reference voltage V.sub.ref remains open. As a result, charge
currently accumulated in each pixel detector element 310 in the
selected pixel array discharges to the readout stage 520, is read
by conversion circuitry 521, and is processed as part of the
current X-ray image being acquired. During implementation of step
1122, accumulated charge present in each pixel detector element 310
in the selected pixel array is reduced from an initial charge value
Q.sup.0 to a remainder charge value Q.sup.Rem at a rate described
by previously presented Equation 1. It is noted that initial charge
value Q.sup.0 is generally different for each pixel detector
element 310. Similarly, remainder charge value Q.sup.Rem is a
function of initial charge value Q.sup.0 and therefore is also
generally different for each pixel detector element 310.
[0069] In step 1123, the controller determines whether there are
remaining pixel arrays to be read out. If yes, method 1100 returns
to step 1121; if no, method 1100 proceeds to step 1131.
[0070] In step 1131, the controller begins reset phase 830. For
example, in some embodiments, the controller causes reset switch
522 to close, and data line 510 is communicatively coupled to
V.sub.ref.
[0071] In step 1132, the controller selects one or more pixel
arrays of X-ray detector 207 to be reset. For example, in the
embodiment of reset phase 830 illustrated in FIG. 10, the
controller selects multiple pixel arrays of X-ray detector 207,
i.e., all M rows of pixel detector elements 310. In such
embodiments, remainder charge value Q.sup.Rem for the pixel
detector elements 310 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 522 during reset phase 830
can be reduced by preventing all pixel arrays of X-ray detector 207
from simultaneously being coupled to data line 510. Thus, in some
embodiments, in step 1132, the controller selects a portion of the
total M rows of pixel arrays to begin being reset in step 1133, 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 FIGS. 12A and 12B.
[0072] FIG. 12A is a timing diagram 1200 schematically illustrating
charge loss for different rows of pixel detector elements 310
during reset phase 1230, according to an embodiment of the present
disclosure. FIG. 12B illustrates a portion of reset phase 1230 for
a first array (row M-4) of pixel detector elements 310 and an
adjacent second array (row M-3) of pixel detector elements 310,
according to an embodiment of the present disclosure.
[0073] Reset phase 1230 is similar to reset phase 830 of FIG. 8, in
that during reset phase 1230, residual charge in each pixel
detector element 310 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 1230, multiple rows of an
X-ray detector panel are undergoing reset simultaneously and
accumulated charge is being transferred simultaneously from all
pixel detector elements 310 associated with the multiple rows.
However, in reset phase 1230, the reset time interval t.sub.reset
begins at different times for different rows of pixel detector
elements 310. Thus, reset of the M rows of pixel detector elements
310 is staged throughout reset phase 1230. Consequently, the larger
initial current in data line 510 that results when a row of pixel
detector elements 310 is initially coupled thereto does not occur
for all rows of pixel detector elements 310 simultaneously.
Instead, the rows of pixel detector elements 310 are initially
coupled to data line 510 and readout stage 522 at different times
in reset phase 1230.
[0074] In the embodiment illustrated in FIGS. 12A and 12B, each of
the M rows of pixel detector elements 310 is initially coupled to
readout stage 522 at a different time in reset phase 1230. Thus, in
such an embodiment, each row of pixel detector elements 310 is
associated with a different reset initiating time. For example, as
shown in FIG. 12B, reset of the pixel detector elements 310 of row
M-4 is initiated at a reset initiating time t.sub.reset, M-4, while
reset of the pixel detector elements 310 of row M-3 is initiated at
a reset initiating time t.sub.reset, M-3. In some embodiments,
t.sub.reset, M-3 is separated in time from reset initiating time
t.sub.reset, M-4 by a staging time interval t.sub.stage. In some
embodiments, staging time interval t.sub.stage is generally uniform
between consecutive rows of pixel detector elements 310, while in
other embodiments, staging time interval t.sub.stage can vary
between different consecutive rows of pixel detector elements 310.
Generally, the duration of staging time interval t.sub.stage
relatively small compared to the duration is of reset phase 1230.
Consequently, in embodiments in which the reset of the M rows of
pixel detector elements 310 is staged throughout reset phase 1230,
the overall duration of reset phase 1230 is not significantly
increased.
[0075] In the embodiment illustrated in FIGS. 12A and 12B, each of
the M rows of pixel detector elements 310 is initially coupled to
readout stage 522 at a unique time in reset phase 1230.
Alternatively, in some embodiments, a first group of multiple rows
of pixel detector elements 310 are initially coupled to readout
stage 522 at one time in reset phase 1230, a second group of
multiple rows of pixel detector elements 310 are initially coupled
to readout stage 522 at a later time in reset phase 1230, a third
group of multiple rows of pixel detector elements 310 are initially
coupled to readout stage 522 at another later time in reset phase
1230, and so on. For example, in one such embodiment, each such
group of multiple rows of pixel detector elements 310 includes 10%
of the M rows of pixel detector elements 310 and each such group
includes different rows of pixel detector elements 310 than the
other rows of pixel detector elements. Thus, in such an embodiment,
reset phase 1230 includes 10 different reset initiating times,
rather than one reset initiating time for each row of pixel
detector elements.
[0076] Returning to FIG. 11, in step 1133, the controller causes
the row or rows of pixel detector elements 310 selected in step
1132 to be reset. For example, in some embodiments, the controller
causes reset switch 522 to close, thereby communicatively coupling
data line 510 to reference voltage V.sub.ref. Furthermore, the
controller causes the readout switch 502 of each pixel detector
element 310 in the selected row or rows to close, thereby
communicatively coupling such pixel detector elements 310 to data
line 510. As a result, charge present in the pixel detector
elements 310 (e.g., remainder charge value Q.sup.Rem) begins to be
transferred to reference voltage V.sub.ref. Thus, the charge
present in the pixel detector elements 310 is reduced at a rate
indicated by previously presented Equation 2.
[0077] In embodiments in which all M rows of pixel detector
elements 310 are selected in step 1131, all pixel detector elements
310 of X-ray imager 207 are reset simultaneously. In such
embodiments, after reset time interval t.sub.reset has elapsed,
method 1100 proceeds to step 1133.
[0078] In embodiments in which a staged reset of pixel arrays is
performed in the reset phase, a single row of pixel detector
elements 310 begins to be reset in step 1132 or a group of rows of
pixel detector elements 310 begins to be reset in step 1132. In
such embodiments, after a staging time interval t.sub.stage has
elapsed, method 1100 proceeds to step 1133.
[0079] In step 1134, the controller determines whether there are
remaining pixel arrays that have not yet begun being reset. If yes,
method returns to step 1132; if no, method 1100 proceeds to step
1135. It is noted that in embodiments in which a staged reset of
pixel arrays is performed in the reset phase, method 1100 may
return to step 1132 while some pixel arrays of X-ray imager 207 are
currently being reset. That is, the controller may perform step
1134 before reset time interval t.sub.reset has elapsed for some or
all pixel arrays that have been selected to be reset.
[0080] In step 1135, the controller determines whether reset has
been completed for all pixel arrays. That is, the controller
determines whether reset time interval t.sub.reset has elapsed for
all pixel arrays of X-ray imager 207. If no, method 1100 returns
back to step 1135; if yes, method 1100 proceeds to step 1136.
[0081] In step 1136, the controller ends the reset phase. For
example, in some embodiments, the controller causes reset switch
522 to open and the readout switch 502 of each pixel detector
element 310 of X-ray imager 207 to open. When acquisition of
further X-ray images is planned, method 1100 returns back to step
1112.
[0082] Implementation of method 1100 enables a significant
reduction in charge remaining in pixel detector elements 310 (for
example from Q.sup.Rem to Q.sup.Fin) over a reset time interval
t.sub.reset. As noted above, reset time interval t.sub.reset 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
t.sub.reset).times.(number of rows of pixel detector elements 310).
Since the number of rows of pixel detector elements 310 can be on
the order of 1000 to 4000, the panel readout time for the
conventional X-ray imager can be significantly increased,
preventing fast image acquisition.
Synchronizing Reset with Application of Treatment Beam
[0083] 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 207 in FIG. 2).
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 830 or reset phase 1230, is
timed to coincide with the application of a treatment beam, such as
treatment beam 230. X-ray imager 207 is insensitive to radiation
during a reset phase as described herein. That is, pixel detector
elements 310 of X-ray imager 207 do not accumulate charge during
such a reset phase, even when treatment beam 230 produces
significant X-ray scatter that is incident on X-ray imager 207. As
a result, the implementation of the reset phase to coincide with a
burst or other application of treatment beam 231 reduces noise that
is normally induced by treatment beam 230. One such embodiment is
illustrated in FIG. 13.
[0084] FIG. 13 is a timing diagram 1300 schematically illustrating
the relative timing of an irradiation phase 1310, a readout phase
1320, and a reset phase 1330 relative to the delivery of a
treatment beam pulse 1350, according to an embodiment of the
present disclosure. As shown, a KV (imaging) beam pulse 1360 occurs
during each irradiation phase 1310 and delivery of treatment beam
pulse 1350 is timed to occur during each reset phase 1330. As
described above, pixel detector elements 310 of X-ray imager 207 do
not accumulate charge during a reset phase. Consequently, even
though scattered X-rays from treatment beam pulses 1350 may strike
the scintillator layer of X-ray imager 207, noise is not added to
the image data generated by X-ray imager 207.
[0085] 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.
[0086] The descriptions of the various embodiments have been
presented for purposes of illustration, but are not intended to be
exhaustive or limited to the embodiments disclosed. Many
modifications and variations will be apparent to those of ordinary
skill in the art without departing from the scope and spirit of the
described embodiments.
[0087] 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.
[0088] Any combination of one or more computer readable medium(s)
may be utilized. The computer readable medium may be a computer
readable signal medium or a computer readable storage medium. A
computer readable storage medium may be, for example, but not
limited to, an electronic, magnetic, optical, electromagnetic,
infrared, or semiconductor system, apparatus, or device, or any
suitable combination of the foregoing. More specific examples (a
non-exhaustive 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. In the context of this document, a computer readable
storage medium may be any tangible medium that can contain, or
store a program for use by or in connection with an instruction
execution system, apparatus, or device.
[0089] While various aspects and embodiments have been disclosed
herein, other aspects and embodiments will be apparent to those
skilled in the art. The various aspects and embodiments disclosed
herein are for purposes of illustration and are not intended to be
limiting, with the true scope and spirit being indicated by the
following claims.
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