U.S. patent application number 16/920034 was filed with the patent office on 2022-01-06 for x-ray ct apparatus with adaptive photon counting detectors.
This patent application is currently assigned to Canon Medical Systems Corporation. The applicant listed for this patent is Canon Medical Systems Corporation. Invention is credited to Liang CAI, Xiaochun LAI, Yi QIANG, Xiaohui ZHAN.
Application Number | 20220000437 16/920034 |
Document ID | / |
Family ID | 1000004955170 |
Filed Date | 2022-01-06 |
United States Patent
Application |
20220000437 |
Kind Code |
A1 |
LAI; Xiaochun ; et
al. |
January 6, 2022 |
X-RAY CT APPARATUS WITH ADAPTIVE PHOTON COUNTING DETECTORS
Abstract
An embodiment of a computed tomography apparatus includes an
x-ray source and scan control circuity configured to control the
x-ray source to expose a subject with x-rays over a scan having a
plurality of views. A detector is disposed to receive x-rays from
the x-ray source, has a plurality of anodes arranged in groups, and
a common conductive strip between the anodes. Photon counting
circuits are respectively provided for each of the anodes and have
adjustable operating parameters. Connection circuitry is configured
to adaptively connect, in a first mode, each anode to one of the
photon counting circuits and, in a second mode, each anode in a
group to a same one of the photon counting circuits. Processing
circuitry, connected to the connection circuitry and the photon
counting circuits, is configured to, for each of the views, select
the first mode or the second mode and adjust the operating
parameters based upon exposure data obtained from exposing the
subject with the x-rays from the x-ray source.
Inventors: |
LAI; Xiaochun; (Vernon
Hills, IL) ; CAI; Liang; (Vernon Hills, IL) ;
QIANG; Yi; (Vernon Hills, IL) ; ZHAN; Xiaohui;
(Vernon Hills, IL) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Canon Medical Systems Corporation |
Otawara-shi |
|
JP |
|
|
Assignee: |
Canon Medical Systems
Corporation
Otawara-shi
JP
|
Family ID: |
1000004955170 |
Appl. No.: |
16/920034 |
Filed: |
July 2, 2020 |
Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61B 6/4241 20130101;
A61B 6/035 20130101; A61B 6/54 20130101 |
International
Class: |
A61B 6/00 20060101
A61B006/00; A61B 6/03 20060101 A61B006/03 |
Claims
1. A computed tomography apparatus, comprising: an x-ray source;
scan control circuity configured to control the x-ray source to
expose a subject with x-rays over a scan having a plurality of
views; a detector disposed to receive x-rays from the x-ray source
and having plurality of anodes arranged in groups; photon counting
circuits respectively provided for each of the anodes and having
adjustable operating parameters; connection circuitry configured to
adaptively connect, in a first mode, each anode to one of the
photon counting circuits and, in a second mode, each anode in a
group to a same one of the photon counting circuits; and processing
circuitry, connected to the connection circuitry and the photon
counting circuits, configured to, for each of the views, select the
first mode or the second mode and adjust the operating parameters
based upon exposure data obtained from exposing the subject with
the x-rays from the x-ray source.
2. The computed tomography apparatus according to claim 1, wherein:
scan control circuity configured to control the x-ray source to
expose the subject with the x-rays over the scan, the scan
including a first exposure before a second exposure; and the
processing circuitry is configured to select the first mode or
second and adjust the operating parameters based upon the first
exposure.
3. The computed tomography apparatus according to claim 2, wherein:
the first exposure has a duration and an x-ray exposure level each
less than that of the second exposure.
4. The computed tomography apparatus according to claim 2, wherein:
each of the photon counting circuits is configured to determine a
number of photons and an energy distribution of the photons for a
corresponding anode during the first exposure.
5. The computed tomography apparatus according to claim 2, wherein:
the processing circuitry comprises: a processor, and a storage
containing mode data and operating parameter data as a function of
x-ray flux and energy; the processor is configured to compare data
from the first exposure of the scan with data in the storage to
select the first mode or the second mode and adjust the operating
parameters; and the scan control circuity is configured to control
the x-ray source to expose the subject with x-rays during the
second exposure of the scan using the selected mode and adjusted
operating parameters from the comparison based upon the data from
first exposure of the scan.
6. The computed tomography apparatus according to claim 5, wherein
the storage is a look up table.
7. The computed tomography apparatus according to claim 1, wherein
the processing circuitry is configured to: generate exposure data;
and process the exposure data to increase spatial uniformity.
8. The computed tomography apparatus according to claim 7, where
the processing circuity is configured to down-sample data collected
from the anodes connected to one of the photon counting circuits
and up-sample data connected from anodes connected in the
group.
9. The computed tomography apparatus according to claim 1, wherein
the photon counting circuitry comprises: a charge sensitive
preamplifier which outputs a pulse signal having a shape and a
width; and shaping circuitry configured to adjust at least one of
the shape and width of the pulse signal.
10. The computed tomography apparatus according to claim 9, wherein
the processing circuitry is configured to adjust at least one of an
integration time of the charge sensitive preamplifier and a pulse
width of the shaping circuitry based upon the exposure data.
11. The computed tomography apparatus according to claim 9, wherein
the shaping circuitry is configured to adjust at least one of the
shape and width of the pulse signal based upon at least one of
x-ray energy and flux.
12. The computed tomography apparatus according to claim 1,
comprising a common conductive strip disposed between the anodes
having a lower applied voltage that that of the anodes.
13. A photon counting computed tomography method, comprising:
exposing a subject with x-rays over a scan having a plurality of
views, each of the views having a first exposure period before a
second exposure period; determining an energy distribution of x-ray
received during the first exposure period using photon counting
circuitry and a detector having a plurality of anodes arranged in
groups, for each of the plurality of views; selecting a number of
anodes in each group and operating parameters of the photon
counting circuitry based upon the x-rays received during the first
exposure period; and exposing the subject during the second period
based upon the selected number of anodes and operating
parameters.
14. The photon counting computed tomography method according to
claim 13, wherein said first exposure period has a duration and an
x-ray exposure level each less than that of the second exposure
period.
15. The photon counting computed tomography method according to
claim 13, comprising: comparing the energy distribution to mode
data and operating parameter data stored as a function of x-ray
flux and energy; and selecting the number of anodes in each group
and the operating parameters of the photon counting circuitry based
upon the comparing.
16. The photon counting computed tomography method according to
claim 13, comprising: exposing the subject to a level of x-rays
during the first period designed to sense x-ray flux and energy of
the view; and exposing the subject to a level of x-rays during the
second period designed to generate image data of the view.
17. The photon counting computed tomography method according to
claim 13, comprising: generating exposure data from exposing the
subject during the second exposure period; and processing the
exposure data to increase spatial uniformity.
18. The photon counting computed tomography method according to
claim 13, comprising: down-sampling data collected from the anodes
connected to one of the photon counting circuitry; and up-sampling
data collected from the anodes connected as a group to the photon
counting circuitry.
19. The photon counting computed tomography method according to
claim 13, wherein the photon counting circuitry comprises: a charge
sensitive preamplifier which outputs a pulse signal having a shape
and a width; and shaping circuitry configured to adjust at least
one of the shape and width of the pulse signal, the method
comprising: adjusting at least one of an integration time of the
charge sensitive preamplifier and a pulse width of the shaping
circuitry based upon the exposure data.
20. The photon counting computed tomography method according to
claim 13, comprising: disposing a common electrode strip between
the anodes in a group; and applying a voltage to the strip less
than that applied to the anodes.
Description
FIELD
[0001] Embodiments described herein relate generally to an X-ray CT
apparatus with photon counting detectors
BACKGROUND
[0002] Conventionally, in medical image systems, such as
photon-counting type X-ray computed tomography (CT) apparatus,
photon-counting type detectors are used to detect X-rays
transmitted through the subject. Photon counting detectors are
designed to record the energy of each incoming X-ray photon. In a
typical CT scan environment, the photon counting detector is
required to be able to record a high flux of incoming photons with
good energy resolving accuracy. Direct-conversion type
semiconductor detectors of cadmium telluride (CdTe), cadmium zinc
telluride (CZT), or the like, or indirect-conversion type detectors
of a scintillator, or the like, are used as a detector. An ASIC is
typically used for the photon-counting CT which amplifies an output
signal from the detector by using an amplifier, shapes its
waveform, and then counts the number of incident X-ray photons of
each of the windows, which are divided in accordance with the level
of the signal.
[0003] The CdTe/CZT sensor is pixelated to a typical size between
250 um to 1 mm. As illustrated in FIG. 1, when an x-ray interacts
within a pixel, induced signal is integrated to a charge sensitive
preamplifier and the output of the charge preamplifier is passed to
shaping, and then signal is compared with different thresholds. The
energy range of x-ray is recorded at counters and arranged in
bins.
[0004] During a CT scan, the flux varies dramatically not only in
the spatial domain and but also time domain. The pixel
configuration (pixel size or pitch size) and circuit parameters,
i.e., decay constants of the preamplifier, shaping constants, and
thresholds of the comparators, etc. significantly affect the
performance of the detector. However, optimized parameters are not
always the same. They are strongly dependent on x-ray flux and
energy.
[0005] For example, a smaller pitch size design (<250 .mu.m
pitch) is preferred for high flux case since it is relatively
immune to pulse-pile up problem. However, this design would at the
same time degrade the detector response as severe charge sharing,
and cross talk effect can happen in such case. The charge loss
between the small pitches is another degradation factor. Larger
pitch size (.about.500 .mu.m) tends to produce less charge sharing
and cross talk effect, leading to better detector response. But the
larger pitch size detector has a strong pule-pile up problem in the
high flux region. Studies have shown that from a material
decomposition noise point of view, smaller pitch size (e.g. <250
.mu.m) is preferred in high flux scenario. However, large pitch
size (e.g. .about.500 .mu.m) is preferred in low flux region.
[0006] The preamplifier parameters also have an effect on the
photon counting. For accurate energy information, it is preferred
to have longer signal integration times for an active reset-type
preamplifier or a longer decay constant for a feedback
resistor-type preamplifier. However, the longer integration time or
decay constant is not preferred when the x-ray flux is high, as
more pile-up will degrade the detector performance. A longer time
for shaping is also preferred for more accurate energy information
but is also subject to pile-up problems. Lastly, optimizing the
thresholds in the signal counting can significantly increase the
signal-to-noise ratio. The optimum thresholds are, however, highly
dependent upon the x-ray flux and spectrum.
[0007] One approach is to use uniform and fixed small pixel
(225-500 .mu.m) pattern design with the readout electronics,
including the preamplifier and shaping electronics, optimized for
high flux. The comparator thresholds are optimized for a specific
case. This design can handle the high flux situation, but
performance is degraded by charge-sharing and cross-talk.
[0008] To address spatial non-uniformity in the x-ray flux, a
detector in U.S. Pat. No. 7,916,836 was proposed having a
non-uniform pixel pattern. In a high flux exposure, the detector is
configured through switching with small pixels (FIG. 6), while in
the low flux exposure, the detector is configured with large pixels
(FIG. 7). However, the x-ray flux varies in both the spatial and
time domains. In one CT scan the same detector can be exposed to
high flux in one view and low flux in another. The flux also can
change scan to scan. The detector in U.S. Pat. No. 7,916,836 is
changed based upon examination protocols or patient size and is not
able to change the pixel pattern in real time to adapt to the
changing flux.
[0009] In U.S. Pat. No. 7,488,945 pixel arrangements can be
adaptively set according to a detected count rate in the previous
views. This allows a detector to adjust the pixel configuration
using the data of the previous view. But such adjusting may not be
fast enough as the flux (count rate) can change dramatically
between two sequential views. Further, the charge loss at streets
between subpixels cannot be avoided. Also, both U.S. Pat. No.
7,916,836 and U.S. Pat. No. 7,488,945 do not adaptively optimizing
operation parameters of readout electronics based on the count rate
and the spectrum. They are critical factors affecting the photon
counting detector and CT performance and also can vary between in
both spatial and time domain.
[0010] Another approach to addressing the high flux and
non-uniformity issues involved a two-layer detector where the first
layer absorbs 90-99% of the flux and the second layer absorbs the
remainder. This system is illustrated in U.S. Pat. No. 7,606,347.
The design involves monitoring subpixels in the first layer,
inhibiting signals from saturated subpixels, and compensating for
the inhibited signals by the second layer detector. The system is
complicated and costly to manufacture.
BRIEF DESCRIPTION OF THE DRAWINGS
[0011] FIG. 1 is a diagram illustrating conventional photon
counting circuitry;
[0012] FIG. 2 is a diagram illustrating an example of the
configuration of a photon-counting type X-ray CT apparatus
according to a first embodiment;
[0013] FIG. 3 is a diagram illustrating an embodiment of the data
acquisition circuitry according to the invention;
[0014] FIG. 4 is a diagram illustrating a modification of the
embodiment of the data acquisition circuitry of FIG. 3;
[0015] FIG. 5 is a diagram illustrating an adjustable photon
counting circuit according to the invention;
[0016] FIG. 6A is a diagram illustrating an example of the charge
sensitive pre-amplifier (CSP) of FIG. 5;
[0017] FIG. 6B is a diagram illustrating another example of the CSP
of FIG. 5;
[0018] FIG. 7A is a diagram illustrating an example of a signal
shaping circuit according to the invention;
[0019] FIG. 7B is a diagram illustrating another example of the
signal shaping circuit according to the invention;
[0020] FIG. 8 is a diagram illustrating an embodiment of the
processing circuit according to the invention;
[0021] FIG. 9 is a flow diagram of the method according to an
embodiment of the invention; and
[0022] FIG. 10 is a flow diagram of post-processing of acquired
data.
DETAILED DESCRIPTION
[0023] With reference to the accompanying drawings, a detailed
explanation is given below of an embodiment of a data acquisition
device and an X-ray CT apparatus. Furthermore, in the following
embodiment, an explanation is given by using, for example, a
photon-counting type X-ray CT apparatus as the X-ray CT
apparatus
First Embodiment
[0024] First, an explanation is given of an embodiment of the
photon-counting type X-ray CT apparatus. FIG. 2 is a diagram that
illustrates an example of the configuration of a photon-counting
type X-ray CT apparatus 1 according to a first embodiment. As
illustrated in FIG. 2, the photon-counting type X-ray CT apparatus
1 according to the first embodiment includes a gantry 10, a bed
device 20, and a console 30.
[0025] The gantry 10 is a device that emits X-rays to a subject P
(patient), detects the X-rays that are transmitted through the
subject P, and outputs them to the console 30, and it includes
X-ray radiation control circuitry 11, an X-ray generation device
12, a detector 13, data acquisition circuitry (DAS: Data
Acquisition System) 14, a rotary frame 15, and gantry drive
circuitry 16.
[0026] The rotary frame 15 is an annular frame that supports the
X-ray generation device 12 and the detector 13 such that they are
opposed to each other with the subject P interposed therebetween
and that is rotated at high speed in a circular orbit around the
subject P by the gantry drive circuitry 16 that is described
later.
[0027] The X-ray radiation control circuitry 11 is a device that
serves as a high-voltage generation unit and supplies a high
voltage to an X-ray tube 12a, and the X-ray tube 12a generates
X-rays by using the high voltage that is supplied from the X-ray
radiation control circuitry 11. Under the control of scan control
circuitry 33, which is described later, the X-ray radiation control
circuitry 11 adjusts the tube voltage or the tube current that is
supplied to the X-ray tube 12a, thereby adjusting the amount of
X-rays that are emitted to the subject P.
[0028] Furthermore, the X-ray radiation control circuitry 11
switches a wedge 12b. Furthermore, the X-ray radiation control
circuitry 11 adjusts the numerical aperture of a collimator 12c,
thereby adjusting the radiation range (the fan angle or the cone
angle) of X-rays. Moreover, according to the present embodiment,
there may be a case where multiple types of wedges are manually
switched by an operator.
[0029] The X-ray generation device 12 is a device that generates
X-rays and emits the generated X-rays to the subject P, and it
includes the X-ray tube 12a, the wedge 12b, and the collimator
12c.
[0030] The X-ray tube 12a is a vacuum tube that emits X-ray beams
to the subject P by using the high voltage that is supplied by the
X-ray radiation control circuitry 11, and it emits X-ray beams to
the subject P in accordance with the rotation of the rotary frame
15. The X-ray tube 12a generates X-ray beams that spread with the
fan angle and the cone angle. For example, under the control of the
X-ray radiation control circuitry 11, the X-ray tube 12a is capable
of continuously emitting X-rays all around the subject P for a full
reconstruction or continuously emitting X-rays for a half
reconstruction within an emission range (180.degree.+the fan angle)
that enables a half reconstruction. Furthermore, under the control
of the X-ray radiation control circuitry 11, the X-ray tube 12a is
capable of intermittently emitting X-rays (pulse X-rays) at a
previously set position (tube position). Furthermore, the X-ray
radiation control circuitry 11 is capable of changing the intensity
of X-rays, emitted from the X-ray tube 12a. For example, the X-ray
radiation control circuitry 11 increases the intensity of X-rays,
emitted from the X-ray tube 12a, at a specific tube position, and
it decreases the intensity of X-rays, emitted from the X-ray tube
12a, in the area other than the specific tube position.
[0031] The wedge 12b is an X-ray filter that adjusts the amount of
X-rays with regard to the X-rays that are emitted from the X-ray
tube 12a. Specifically, the wedge 12b is a filter that transmits
and attenuates X-rays, emitted from the X-ray tube 12a, such that
X-rays, emitted from the X-ray tube 12a to the subject P, has a
predetermined distribution. For example, the wedge 12b is a filter
that is obtained by processing aluminum so as to have a
predetermined target angle or a predetermined thickness.
Furthermore, the wedge is also called a wedge filter or a bow-tie
filter.
[0032] The collimator 12c is a slit that narrows the irradiation
range of X-rays, of which the amount of X-rays has been adjusted by
the wedge 12b, under the control of the X-ray radiation control
circuitry 11 that is described later.
[0033] The gantry drive circuitry 16 drives and rotates the rotary
frame 15 so that the X-ray generation device 12 and the detector 13
are rotated in a circular orbit around the subject P.
[0034] Each time an X-ray photon enters, the detector 13 outputs
the signal with which the energy value of the X-ray photon may be
measured. The X-ray photon is, for example, an X-ray photon that is
emitted from the X-ray tube 12a and is transmitted through the
subject P. The detector 13 includes multiple detection elements
that output an electric signal (analog signal) of 1 pulse each time
an X-ray photon enters. The photon-counting type X-ray CT apparatus
1 counts the number of electric signals (pulses) so as to count the
number of X-ray photons that enter each of the detection elements.
Furthermore, the photon-counting type X-ray CT apparatus 1 performs
arithmetic processing on the signal so as to measure the energy
value of the X-ray photon that causes output of the signal.
[0035] The above-described detection element includes, for example,
a scintillator and an optical sensor, such as a photomultiplier
tube. In such a case, the detector 13, illustrated in FIG. 2, is an
indirect-conversion type detector that converts the incident X-ray
photon into scintillator light by using the scintillator and
converts the scintillator light into an electric signal by using
the optical sensor, such as a photomultiplier tube. Furthermore,
there may be a case where the above-described detection element is
a semiconductor device of, for example, cadmium telluride (CdTe),
cadmium zinc telluride (CdZnTe), or the like. In such a case, the
detector 13, illustrated in FIG. 2, is a direct-conversion type
detector that directly converts the incident X-ray photon into an
electric signal.
[0036] For example, the detector 13, illustrated in FIG. 2, is a
plane detector in which detection elements are arranged in N
columns in the channel direction (the direction of the X axis in
FIG. 2) and in M columns in the direction of the rotational center
axis of the rotary frame 15 (the direction of the Z axis in FIG. 2)
where the gantry 10 is not tilted. When a photon enters, the
detection element outputs an electric signal of 1 pulse. The
photon-counting type X-ray CT apparatus 1 discriminates among
individual pulses that are output from a detection element 131,
thereby counting the number of X-ray photons that enter the
detection element 131. Furthermore, the photon-counting type X-ray
CT apparatus 1 performs arithmetic processing based on the
intensity of a pulse, thereby measuring the energy value of the
counted X-ray photon.
[0037] The data acquisition circuitry 14 is a data acquisition
system (DAS), and it acquires the detection data on X-rays that are
detected by the detector 13. For example, the data acquisition
circuitry 14 generates the count data that is obtained by counting
the photons (X-ray photons), which come from the X-ray that is
transmitted through the subject, for each energy band, and it
transmits the generated count data to the console 30 that is
described later. For example, if X-rays are continuously emitted
from the X-ray tube 12a while the rotary frame 15 is rotated, the
data acquisition circuitry 14 acquires the group of count data for
the entire periphery (360 degrees). The data acquisition circuitry
14 also can acquire data for each view. Furthermore, the data
acquisition circuitry 14 transmits each acquired count data in
relation to the tube position to the console 30 that is described
later. The tube position is the information that indicates the
projection direction of the count data.
[0038] The bed device 20 is a device on which the subject P is
placed and, as illustrated in FIG. 2, it includes a bed drive
device 21 and a top board 22. The bed drive device 21 moves the top
board 22 in the direction of the Z axis to move the subject P into
the rotary frame 15. The top board 22 is a board on which the
subject P is placed. Furthermore, in the present embodiment, an
explanation is given of a case where the relative position between
the gantry 10 and the top board 22 is changed by controlling the
top board 22; however, this is not a limitation on the embodiment.
For example, if gantry 10 is self-propelling, the relative position
between the gantry 10 and the top board 22 may be changed by
controlling driving of the gantry 10.
[0039] Furthermore, for example, the gantry 10 conducts helical
scan to scan the subject P in a helical fashion by rotating the
rotary frame 15 while the top board 22 is moved. Alternatively, the
gantry 10 conducts conventional scan to scan the subject P in a
circular orbit by rotating the rotary frame 15 with the position of
the subject P fixed after the top board 22 is moved. Alternatively,
the gantry 10 implements a step-and-shoot method to conduct
conventional scan at multiple scan areas by moving the position of
the top board 22 at a constant interval.
[0040] The console 30 is a device that receives an operation of the
photon-counting type X-ray CT apparatus 1 from an operator and that
reconstructs X-ray CT image data by using the projection data that
is acquired by the gantry 10. As illustrated in FIG. 2, the console
30 includes input circuitry 31, a display 32, the scan control
circuitry 33, preprocessing circuitry 34, memory circuitry 35,
image reconstruction circuitry 36, and processing circuitry 37.
[0041] The input circuitry 31 includes a mouse, keyboard,
trackball, switch, button, joystick, or the like, which is used by
an operator of the photon-counting type X-ray CT apparatus 1 to
input various commands or various settings, and it transfers the
information on the command or setting, received from the operator,
to the processing circuitry 37. For example, the input circuitry 31
receives, from an operator, a capturing condition for X-ray CT
image data, a reconstruction condition for reconstructing X-ray CT
image data, an image processing condition for X-ray CT image data,
or the like.
[0042] The display 32 is a monitor that is viewed by an operator
and, under the control of the processing circuitry 37, it displays
the image data, generated from X-ray CT image data, to the operator
or displays a graphical user interface (GUI) for receiving various
commands, various settings, or the like, from the operator via the
input circuitry 31.
[0043] The scan control circuitry 33 controls operations of X-ray
radiation control circuitry 11, the gantry drive circuitry 16, the
data acquisition circuitry 14, and the bed drive device 21 under
the control of the processing circuitry 37, thereby controlling
data acquisition processing by the gantry 10. For example, scan
control circuitry 33 sends sequence control commands to data
acquisition circuitry 14 to control exposure operations, as
discussed in more detail below.
[0044] The preprocessing circuitry 34 performs correction
processing, such as logarithmic conversion processing, offset
correction, sensitivity correction, or beam hardening correction,
on the count data that is generated by the data acquisition
circuitry 14, thereby generating corrected projection data.
[0045] The memory circuitry 35 stores the projection data that is
generated by the preprocessing circuitry 34. Furthermore, the
memory circuitry 35 stores the image data, or the like, which is
generated by the image reconstruction circuitry 36 that is
described later. Moreover, the memory circuitry 35 appropriately
stores processing results of the processing circuitry 37 that is
described later.
[0046] The image reconstruction circuitry 36 reconstructs X-ray CT
image data by using the projection data that is stored in the
memory circuitry 35. Here, the reconstruction method includes
various methods, and it may be, for example, back projection
processing. Furthermore, the back projection processing may
include, for example, back projection processing by using a
filtered back projection (FBP) method. Alternatively, the image
reconstruction circuitry 36 may also use a successive approximation
technique to reconstruct X-ray CT image data. Furthermore, the
image reconstruction circuitry 36 conducts various types of image
processing on X-ray CT image data, thereby generating image data.
Then, the image reconstruction circuitry 36 stores, in the memory
circuitry 35, the reconstructed X-ray CT image data or the image
data that is generated during various types of image
processing.
[0047] The processing circuitry 37 controls operations of the
gantry 10, the bed device 20, and the console 30 so as to perform
the overall control on the photon-counting type X-ray CT apparatus
1. Specifically, the processing circuitry 37 controls the scan
control circuitry 33 so as to control CT scan that is conducted by
the gantry 10. Furthermore, the processing circuitry 37 controls
the image reconstruction circuitry 36 so as to control image
reconstruction processing or image generation processing by the
console 30. Furthermore, the processing circuitry 37 performs
control such that various types of image data, stored in the memory
circuitry 35, are displayed on the display 32.
[0048] Heretofore, the overall configuration of the photon-counting
type X-ray CT apparatus 1 according to the first embodiment is
explained. Here, each processing function, performed by each of the
above-described circuitry, is stored in the memory circuitry 35 in
the form of the program that is executable by the computer.
Furthermore, each circuitry reads and executes each program from
the memory circuitry 35, thereby performing the above-described
various functions.
[0049] In one example, programs corresponding to the operations of
the data acquisition circuitry 14 are stored in the memory
circuitry 35 in the form of a program that is executable by a
computer. Processor 37 executes the programs for data acquisition
circuitry 14 and sends instructions to and controls data
acquisition circuitry 14 to acquire data as well as controls the
transfer data from data acquisition circuitry 14. In a second
example, data acquisition circuitry 14 includes a processor that
reads and executes each program from the memory circuitry 35 to
implement the function that corresponds to each program.
[0050] Furthermore, the word "processor", used in the above
explanations, means for example a central processing unit (CPU), a
graphics processing unit (GPU), or a circuit, such as an
application specific integrated circuit (ASIC), or a programmable
logic device (e.g., a simple programmable logic device: SPLD, a
complex programmable logic device: CPLD, or a field programmable
gate array: FPGA). The processor reads and executes the program,
stored in the memory circuitry, to perform the function.
Furthermore, a configuration may be such that, instead of storing a
program in the memory circuitry, a program is directly installed in
a circuit of the processor. In this case, the processor reads and
executes the program, installed in the circuit, to perform the
function. Furthermore, with regard to the processors according to
the present embodiment, instead of the case where each processor is
configured as a single circuit, multiple independent circuits may
be combined to be configured as a single processor to implement the
function.
[0051] With the above-described configuration, the photon-counting
type X-ray CT apparatus 1 according to the first embodiment allows
an improvement in the image acquisition due to the operation of the
data acquisition circuitry 14 that optimizes data acquisition based
upon energy and flux, which is described in detail below. A diagram
of the data acquisition circuitry 14 is provided in FIG. 3. A group
of anodes in the detector is shown at 40. The anodes 40-1 to 40-4
have a typical size of 500 um or less. These anodes are selectively
grouped to form one large pixel and are individually used as plural
small pixels. In this example, four adjacent anodes 40-1 to 40-4
arranged in a rectangular group are combined to form a large pixel
and each individual anode forms a small pixel. Other arrangements
and numbers of anodes could be grouped together to form the large
pixel. For example, 1.times.2, 1.times.4 or 3.times.3 groups of
anodes are possible.
[0052] A street is formed between the anodes and is typically 10 um
to 50 um wide. In the center of the street, there is common
conductive strip 40-5 with a typical size of 5 um to 45 um, which
has lower voltage compared to anodes 40-1 to 40-4. It prevents the
charge generated by x-ray from being trapped in the street. The
anode is typically grounded to 0 V and the strip is typically
biased in the range of 0V to -100V.
[0053] The outputs of the anodes are connected to switching circuit
41. Switching circuit 41 connects one of anodes 40-1 to 40-4 to a
respective one of the adjustable photons counting circuits 42-1 to
42-4 when the switches 44 are in the upper position. The data from
each anode is collected individually by adjustable photon counting
circuits 42-1 to 42-4. Switching circuit 41 connects all of the
anodes to adjustable photon counting circuit 42-5 when the switches
44 are in the lower position, aggregating the individual anodes to
one large pixel where the data is collected together.
[0054] The adjustable photon counting circuits 42-1 to 42-5 count
the photons and collect the photon counts into energy bin to
generate collected data. The collected data is output to processing
circuit 43.
[0055] Processing circuit 43 receives, analyzes and stores, in
storage 45, the collected data in both of sense and exposure
operations. Storage 45 may be any type of memory such as a RAM or
NAND. The processing circuit receives data from the sensing
operation to determine the flux of the x-ray signal and generate a
switching signal and one or more parameter signals for optimal data
collection. After the sensing operation, the exposure operation
takes place where the anodes are optimally configured as a small or
large pixel and the adjustable photon counting circuits 42-1 to
42-5 have optimal settings to collect the exposure data.
[0056] Processing circuit 43 generates the switching signal to
control switching circuit 41 to appropriately connect the anodes
40-1 to 40-4 as individual (small) pixels or one large pixel, and
generates one or more parameter signals to adjust the operation of
adjustable photon counting circuits 42-1 to 42-5. It is noted that
while the parameter signal line is shown as a single contact with
circuits 42-1 to 42-4, each of circuits 42-1 to 42-4 is connected
to the parameter signal line to receive the parameter signal or
signals. The processing 43 will be described in more detail
below.
[0057] A modification of the circuit of FIG. 3 is shown in FIG. 4.
There are only four adjustable photon counting electronics 42-1 to
42-4 and circuit 42-4 collects both the data from anode 40-4 when
the circuit is connected in individual anode mode (small size
pixel) and the data from all of anodes 40-1 to 40-4 when the
circuit is connected in the large pixel mode. This modification has
fewer adjustable photon counting circuits resulting in a simpler
and less expensive design.
[0058] FIG. 5 shows the adjustable photon counting electronics 42-1
to 42-5 in more detail. A charge sensitive preamplifier (CSP) 46
(Details are discussed below with respect to FIGS. 6A and 6B)
receives the current output of one of the pixels generated by a
photon and this current is accumulated in the feedback capacitors
(Cp) of CSP and converted to a voltage signal output. The signal
output by CSP 46 is input to shaping circuit 47 which turns the
signal to desired shapes, typically with a reduced pulse width but
the preserved energy information of the signal. The desired shapes
are discussed below with respect to FIGS. 7A and 7B.
[0059] The signal output from the shaping circuit 47 is input to a
comparator 48 having a plurality of comparing circuits each
comparing the shaped signal to a reference voltage generated by
reference voltage circuits V1-V5. The reference voltages are chosen
to arrange the number of counts into energy bins representing the
energy of the signal output by shaping circuit 47. In this
embodiment, five comparing circuits compare the signal output by
shaping circuit 47 to reference voltages of reference voltage
circuits V1-V5, respectively, and the five counters record the
number of times the shaping signal has a voltage value in the
respective energy bins. The present invention is not limited to 5
comparing circuits. Other numbers of comparing and reference
voltage circuits may be used based upon system requirements and
design.
[0060] The parameter signals may also include a setting value for
the reference voltage circuits V1-V5. The voltages of the reference
voltage circuits V1-V5 are optimally set to collect the data based
upon the results of the sensing operation.
[0061] FIGS. 6A and 6B show two examples of the CSP circuit 46.
FIG. 6A shows an active reset type CSP and FIG. 6B shows a feedback
resistor type CSP. To get accurate energy information, it is
preferred to have a longer signal integration time. On the other
hand, longer integration time is not preferable when x-ray flux is
high, given that more pile-up will degrade the detector
performance.
[0062] In FIG. 6A, the switch is open during the charge integration
period and the switch 50 is closed to reset amplifier 51. The
appropriate time is in the range of 5 ns to 100 ns and is chosen
based upon the flux and energy readings from the sensing operation.
For FIG. 6B, the value of variable resistor 52 is chosen to set the
appropriate decay constant (R*Cp), which is in the range of 5 ns to
100 ns. Sensing circuit 43 outputs parameter signals to
appropriately set the integration time or decay constant. The
present invention is not limited to these two types preamplifiers
and other types of preamplifiers, such as periodically reset
preamplifiers, may be used depending upon system requirements and
design.
[0063] FIG. 7A and FIG. 7B show two examples of shaping circuit 46.
FIG. 7A shows an adjustable CR-RC shaper and FIG. 7B shows an
adjustable single-line delay shaper. To get accurate energy
information, it is preferred to have a wider pulse width (longer
shaping time). On the other hand, wider pulse width is not
preferable when x-ray flux is high, given that more pile-up will
degrade the detector performance.
[0064] In FIG. 7A, the value of the variable resistor (R.sub.1 and
R.sub.2) is chosen to set an appropriate shaping width, which is
determined by Cp.sub.1*R.sub.1 and Cp.sub.2*R.sub.2, based on the
flux and energy reading from the sensing operation. E.sub.out and E
are the signals before and after the shaping respectively, and are
given as
E out = EC p .times. 1 .times. R 1 C p .times. 1 .times. R 1 - C p
.times. 1 .times. R 1 .function. [ e - t C p .times. 1 .times. R 1
- e - t C p .times. 2 .times. R 2 ] , ##EQU00001##
where t is time.
[0065] Similarly, in FIG. 7B, the value of the variable delay time
T is chosen to be set appropriately based on the flux and energy
reading from the sensing operation. The present invention is not
limited to these two types preamplifiers and other types of
preamplifiers, such as a multiple stage CR-RC.sup.n shaper or
multiple line delay shaper, active pulse shaper, triangular shaper
and trapezoidal shaper, may be used based upon system requirements
and design. The key feature of the shaping circuit according to the
invention is that the shaping parameters of these shapers can be
adjusted based input x-ray flux and energy.
[0066] FIGS. 7A and 7B each show the original signal and two shaped
signals. The longer delay time preserves the signal shape better
but has a wider pulse width. The shorter delay time has the
advantage of a shorter pulse width but preserves the signal
worse.
[0067] The processing circuit 43 is shown in more detail in FIG. 8.
Circuit 43 includes a processor 53 connected to a set of registers
52 and a LUT 54. The LUT stores optimal settings for the anodes,
CSP, signal shaping circuit, and reference voltage levels as a
function of energy and flux values.
[0068] In one embodiment, the optimal settings are pre-calculated
by the following method:
x ^ = argmin x .times. { .PHI. .function. ( x | f ) }
##EQU00002##
where {circumflex over (x)}=[{circumflex over (x)}.sub.1,
{circumflex over (x)}.sub.2, {circumflex over (x)}.sub.3,
{circumflex over (x)}.sub.4] is the optimized settings for the
anodes ({circumflex over (x)}.sub.1), CSP ({circumflex over
(x)}.sub.2), signal shaping ({circumflex over (x)}.sub.3) and the
comparators' thresholds ({circumflex over (x)}.sub.4). It will
minimize the object function .PHI.(x|f) with given input,
(f=[f.sub.1, f.sub.2]) where f.sub.1 is x-ray flux and f.sub.2, a
metric indicating whether the spectrum is soft or hard, such as
ratio between the number of low and high energy x-rays. The object
function (.PHI.(x|f)) can be specifically designed base on system
and design requirement, such as material decomposition variance,
material decomposition bias, or mean square error.
[0069] The precalculated optimized parameters {circumflex over
(x)}=[{circumflex over (x)}.sub.1, {circumflex over (x)}.sub.2,
{circumflex over (x)}.sub.3, {circumflex over (x)}.sub.4], could be
saved in LUT. Table 1 gives an example of LUT generated by
minimizing the decomposition variance:
.PHI.(x|f)=.SIGMA..sub.i.sup.2m.sub.ii, M=1/J(x|f), where J(x|f) is
the Fisher information matrix of decomposed material length and
m.sub.ii is the diagonal element of matrix M. In this table, two
flux scenarios (low fluxes and high fluxes) and two spectrum
scenarios (soft and hard) are given, with total four cases. Please
note that the LUT is not limited to only have four cases, and could
have multiple scenarios for fluxes and spectrum.
TABLE-US-00001 TABLE 1 LUT used for switching the anode
configuration and circuitry operation conditions Soft
spectrum.sup.2 Hard spectrum f.sub.2 .ltoreq. F.sub.2 f.sub.2 >
F.sub.2 Low flux.sup.1 Pixel Size.sup.3: {circumflex over
(x)}.sub.1 = 1 Pixel Size: {circumflex over (x)}.sub.1 = 1 f.sub.1
.ltoreq. F.sub.1 CSP decay or reset CSP decay or reset time.sup.4:
time: ({circumflex over (x)}.sub.2 = T.sub.11) ({circumflex over
(x)}.sub.2 = T.sub.12) Shaping Time.sup.5: Shaping Time:
({circumflex over (x)}.sub.3 = S.sub.11) ({circumflex over
(x)}.sub.3 = S.sub.12) Comparator Voltage Comparator Voltage
thresholds.sup.6: ({circumflex over (x)}.sub.4 = V.sub.11)
thresholds: ({circumflex over (x)}.sub.4 = V.sub.12) High flux:
Pixel Size: {circumflex over (x)}.sub.1 = 0 Pixel Size: {circumflex
over (x)}.sub.1 = 0 f.sub.1 > F.sub.1 CSP decay or reset CSP
decay or reset time: time: ({circumflex over (x)}.sub.2 = T.sub.21)
({circumflex over (x)}.sub.2 = T.sub.22) Shaping Time: Shaping
Time: ({circumflex over (x)}.sub.3 = S.sub.21) ({circumflex over
(x)}.sub.3 = S.sub.22) Comparator Voltage Comparator Voltage
thresholds: thresholds: ({circumflex over (x)}.sub.4 = V.sub.21)
({circumflex over (x)}.sub.4 = V.sub.22) Notes: .sup.1Threshold
F.sub.1 is used to define the high flux and low flux case,
typically in the range of 1 Mcps/mm.sup.2 to 200 Mcps/mm.sup.2.
.sup.2Threshold F.sub.2 is used to determine whether the spectrum
is soft or hard. It can use ratio between the number of low and
high energy x-rays, typically in the range of 0.1 to 10.
.sup.3Pixel size {circumflex over (x)}.sub.1 = 1 indicates that the
circuitry is arranged with the large anode pixel and 0, with the
small pixel. .sup.4CSP decay or reset time, {circumflex over
(x)}.sub.2 is in the range of 5 ns to 100 ns; typically, T.sub.21
< T.sub.22 < T.sub.11 < T.sub.12 .sup.5Shaping time,
{circumflex over (x)}.sub.3 is in the range of 5 ns to 100 ns;
typically, S.sub.21 < S.sub.22 < S.sub.11 < S.sub.12
.sup.6Comparator voltage thresholds, {circumflex over (x)}.sub.4,
correspond to the photon energy within 0 keV to 160 keV; typically,
V.sub.11 < V.sub.21 < V.sub.12 < V.sub.22
[0070] Processor 53 controls the registers and outputs the
switching signal and parameter signals based upon the contents of
the registers and the information in the LUT. In a first operation,
processor 53 receives an instruction from the sequence control
(i.e., scan control circuit 33) that imaging for a view is ready
and then resets the registers. Next, after a predetermined delay,
the system performs the sensing operation to expose the subject to
the sensing x-rays. Data from the adjustable photon counting
circuits 42-1 to 42-n are stored in the set of registers. The
processor compares the data in the registers with the information
in the LUT and outputs a switching signal to switch 41 and
parameter signals to the adjustable photon counting circuits 42-1
to 42-n such as the integration or delay time of the
pre-amplifiers, shaping circuit parameters and reference voltage
levels. For example, if the flux values are larger than a preset
threshold value, the processor outputs a switching signal to
configure the anodes in the small pixel mode, CSP using a shorter
decay or reset time, filter/shaper with a shorter shaping time, and
comparators with higher thresholds settings that can capture pile
up effects better. Conversely, if the flux values are lower the
preset threshold value, then the processor outputs a switching
signal to configure the anodes together in the large pixel mode,
CSP using a longer decay or reset time, filter/shaper with a longer
shaping time, and comparators with lower thresholds settings. The
threshold value will depend on the size of the anodes and the
number anodes in a group. Specific parameters' value for CSP,
filter and comparator thresholds are typically in the range of 5 ns
to 100 ns (CSP), 5 ns to 100 ns (Shaper), 0 keV to 160 keV
(comparator thresholds), respectively and will be precalculated and
dependent on the object function (.PHI.(x|f)), which can be
specifically designed base on system and design requirement, such
as material decomposition variance, material decomposition bias, or
mean square error.
[0071] Once the switch 41 and adjustable photon counting
electronics 42-1 to 42-n are optimally set, the sequence control
instructs the system to expose the subject for a given time with
selected scan parameters over the view and collect exposure data.
Once the exposure is completed, the processor instructs the
registers to output the data.
[0072] The system according to the first embodiment optimally
arranges the anodes spatially over the entire detector. The
arrangement can have both small and large pixels configured to
optimally acquire data from the x-ray exposure. Similarly, the
associated settings for the photon counting circuits, shaping
circuits and reference voltages are also arranged spatially over
the entire detector. The detector can be spatially optimized to
collect exposure data.
[0073] In a modification of the above embodiment, instead of each
sensing circuit having a processor and LUT, the circuitry can be
arranged to have one central processor with an LUT that processes
the sensing data from all of the registers. As a second
modification, several processors may be arranged to process a given
number of registers. These two modifications have less
circuitry.
[0074] The detector has non-uniform spatial sampling. The data can
be processed to compensate for the non-uniform spatial sampling to
achieve uniform sampling. In one approach, the large pixel data can
be up-sampled to the high resolution of the small pixel. In a
second approach, the small pixel data will be combined to the lower
resolution of the large pixel through down-sampling. A third
approach is to use a reconstruction algorithm to process the
non-uniform sampling. This reconstruction algorithm can be typical
iteratively reconstruction method and it can reconstruct the
imaging object which is uniformly sampled from the non-uniform
detector sampling data. The up-sampling, down-sampling and
reconstruction algorithm are performed in the preprocessing
circuitry 34.
[0075] The system according to the invention provides several
advantages. Compared with conventional uniform pixel pitch design,
the adaptive pixel configuration maintains good information in both
high and low flux scenario, with ignorable dose penalty. The
adaptive pixel configuration also allows the detector to always
work in an optimal mode, improving the information collection
efficiency, and detector response while reducing overall radiation
dose and pile-up problems. Further, the overall power consumption
can be reduced because the low flux regions are switched to the
large pixel mode and the total number of active channels are
reduced. This will reduce thermal the management effort in photon
counting CT, which is always one of the major hurdles to develop
efficient photon counting CT.
[0076] In a second embodiment of the system, different photon
counting circuitry may be used. The system is not limited to the
embodiment above using pre-amplifiers, shaping circuitry and
voltage comparators. The adaptive pixel configuration may be used
to optimize the performance of such other photon counting circuitry
designed without shaping circuitry or without pixel switching
circuitry. Generally speaking, leaving out shaping or switching
will degrade the detector performance but may simplifying the
circuitry.
[0077] The switching circuits 41, adjustable photon counting
circuits 42-n and the processing circuit 43 for the entire detector
are preferably implemented as an ASIC. Other configurations using
separate circuit components are also possible.
[0078] A first embodiment of the method according to the invention
is shown in FIG. 9. A user operates the input circuitry by, for
example, entering commands, instructions, and/or parameters via a
keyboard or touch screen (step 60). The CT apparatus generates an
exposure plan to scan a subject over a plurality of views in step
61. For each view beginning with the first view (step 62), in a
first period the CT apparatus conducts a sensing operation to sense
the flux and energy of the x-ray. The subject is exposed for a
short amount of time to minimize exposure dose, such as a few tens
of microseconds, the dose of which is, for example, <1% of dose
used to generate the exposure data. The detector captures the
sensing exposure data (flux and energy) and the photon counting
circuits generates the collected data in the energy bins (step
63).
[0079] From the sensing exposure data the CT apparatus, in a second
period, sets an optimal photon-counting mode. The setting of
optimal mode includes setting an optimal pixel configuration over
the entire detector geometry by spatially selecting optimal pixel
sizes, and spatially setting optimal pre-amplifier, shaping and
energy comparator threshold settings over the entire detector
geometry (step 64). One or more of these parameters are adjusted to
obtain optimal exposure acquisition. The selection process can be
achieved using a look-up table (LUT) populated with predetermined
settings derived empirically or derived experimentally through a
calibration process using known samples and/or sample materials.
The second period typically in the range of 1 ns to 1 .mu.s.
[0080] In a third period, the subject is exposed over the view and
the exposure data is acquired (step 65). The acquisition typically
lasts on the order of hundreds to thousands of microseconds. As
such, the dose penalty for the sensing and parameter setting is
only about 1%. The exposure data
[0081] The sensing, parameter setting and data acquisition
procedures are repeated for each view until the last view is
reached (step 66) and the procedure ends (step 67).
[0082] FIG. 10 shows further steps of the method according to the
invention. After acquiring the exposure data in step 65 for all of
the views, the user of the system can enter an instruction for the
acquired data to be post-processed (step 70). The acquired data is
processed by up-sampling (step 71), down-sampling (step 72) or by
running a reconstruction algorithm (step 73) according to the
instruction. Post-processed data is obtained (step 74).
* * * * *