U.S. patent application number 16/149916 was filed with the patent office on 2019-01-31 for photon counting apparatus.
This patent application is currently assigned to Toshiba Medical Systems Corporation. The applicant listed for this patent is Toshiba Medical Systems Corporation. Invention is credited to Emi Tamura.
Application Number | 20190029622 16/149916 |
Document ID | / |
Family ID | 55631907 |
Filed Date | 2019-01-31 |
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United States Patent
Application |
20190029622 |
Kind Code |
A1 |
Tamura; Emi |
January 31, 2019 |
PHOTON COUNTING APPARATUS
Abstract
According to an embodiment, The X-ray tube generates X-rays. The
X-ray detector detects the X-rays transmitted through a subject.
The data acquisition circuitry acquires count data concerning a
count number of the detected X-rays for energy bands. The memory
circuitry stores data of a response function that associates
incident X-rays on the X-ray detector with a response
characteristic of a system including the X-ray detector and the
data acquisition circuitry. The processing circuitry calculates an
X-ray absorption amount of each of a plurality of base substances
based on the count data concerning the energy bands acquired by the
data acquisition circuitry, an energy spectrum of the incident
X-rays, and the response function.
Inventors: |
Tamura; Emi; (Nasushiobara,
JP) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Toshiba Medical Systems Corporation |
Otawara-shi |
|
JP |
|
|
Assignee: |
Toshiba Medical Systems
Corporation
Otawara-shi
JP
|
Family ID: |
55631907 |
Appl. No.: |
16/149916 |
Filed: |
October 2, 2018 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
14870386 |
Sep 30, 2015 |
10117628 |
|
|
16149916 |
|
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|
Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61B 6/032 20130101;
A61B 6/482 20130101; A61B 6/42 20130101; A61B 6/5205 20130101; A61B
6/4035 20130101; A61B 6/4233 20130101; A61B 6/4241 20130101 |
International
Class: |
A61B 6/00 20060101
A61B006/00; A61B 6/03 20060101 A61B006/03 |
Foreign Application Data
Date |
Code |
Application Number |
Oct 1, 2014 |
JP |
2014-203368 |
Sep 29, 2015 |
JP |
2015-191306 |
Claims
1. A photon counting apparatus, comprising: an X-ray to be
configured to generate X-rays; an X-ray detector configured to
detect the X-rays generated by the X-ray tube and transmitted
through a subject; data acquisition circuitry configured to acquire
count data concerning a count number of the detected X-rays for at
least 16 energy bands, based on an output signal from the X-ray
detector; memory circuitry configured to store data of a response
function that associates an energy of incident X-rays on the X-ray
detector with a response characteristic of a system including both
of the X-ray detector and the data acquisition circuitry; and
processing circuitry configured to calculate an X-ray absorption
amount of each of a plurality of base substances based on the count
data concerning the at least 16 energy bands acquired by the data
acquisition circuitry, an energy spectrum of the incident X-rays
generated by the X-ray tube, and the data of the response function
stored in the memory circuitry.
2. An image generating apparatus, comprising: memory circuitry
configured to store data of a response function that associates
energy of incident X-rays on an X-ray detector with a response
characteristic of a system including both of the X-ray detector and
data acquisition circuitry; and processing circuitry configured to
acquire count data concerning a count number of X-rays for at least
16 energy bands; and calculate an X-ray absorption amount of each
of a plurality of base substances based on the count data
concerning the at least 16 energy bands acquired by the data
acquisition circuitry, an energy spectrum of the incident X-rays
generated by the X-ray tube, and the data of the response function
stored in the memory circuitry.
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This application is a continuation Application of U.S.
application Ser. No. 14/870,386, filed Sep. 30, 2015, which is
based upon and claims the benefit of priority from the prior
Japanese Patent Application No. 2014-203368, filed Oct. 1, 2014 and
the prior Japanese Patent Application No. 2015-191306, filed Sep.
29, 2015, the entire contents of all of which are incorporated
herein by reference.
FIELD
[0002] Embodiments described herein relate generally to a photon
counting apparatus.
BACKGROUND
[0003] A photon counting apparatus such as a photon counting CT
apparatus detects X-rays of a high dose on a photon basis and
discriminates a substance that the X-rays have passed through. In
the photon counting apparatus, a direct type detector such as a
semiconductor detector having an excellent energy resolving power
is used to discriminate a substance. As a readout circuitry used to
read out a signal from the detector, a highly integrated element
such as an ASIC is assumed to be used. Regardless of its high
energy resolving power, the semiconductor detector has many
problems such as a high cost and poor stability when used in the
photon counting apparatus. On the other hand, an X-ray computed
tomography apparatus uses an indirect type detector. The indirect
type detector is formed by combining a phosphor such as a
scintillator and a photodetector such as a photomultiplier. Such an
indirect type detector is field-proven as a detector tor an X-ray
CT and features a low cost and high stability, as compared to the
semiconductor detector. However, because of a low energy resolving
power, it is very difficult to discriminate a substance.
BRIEF DESCRIPTION OF THE DRAWING
[0004] FIG. 1 is a block diagram shewing the arrangement of a
photon counting CT apparatus according to the embodiment.
[0005] FIG. 2 is a graph showing an energy spectrum I.sub.0(E) of
X-rays generated by an X-ray tube without a wedge filter;
[0006] FIG. 3 is a graph showing the output of a standard detection
system using LaBr3 as a scintillator with respect to monochromatic
X-rays;
[0007] FIG. 4 is a graph showing the output of the standard
detection system using LaBr3 as a scintillator with respect to
other monochromatic X-rays;
[0008] FIG. 5 is a graph schematically showing a response function
according to the embodiment;
[0009] FIG. 6 is a graph showing comparison between an energy
spectrum represented by count data of actual measurement and an
energy spectrum represented by a model function;
[0010] FIG. 7 is a block diagram showing an example of the
arrangement of a data acquisition circuitry according to the
embodiment;
[0011] FIG. 8 is a block diagram showing another example of the
arrangement of the data acquisition circuitry according to the
embodiment; and
[0012] FIG. 9 is a flowchart showing the typical procedure of
photon counting CT imaging performed under the control of a system
control circuitry shown in FIG. 1.
DETAILED DESCRIPTION
[0013] In general, according to one embodiment, a photon counting
apparatus includes an X-ray tube, an X-ray detector, a data
acquisition circuitry, a memory circuitry, and a processing
circuitry. The X-ray tube generates X-rays. The X-ray detector
detects the X-rays generated by the X-ray tube and transmitted
through a subject. The data acquisition circuitry acquires count
data concerning a count number of the detected X-rays for a
plurality of energy band; based on an output signal from the X-ray
detector. The memory circuitry stores data of a response function
that associates incident X-rays on the X-ray detector with a
response characteristic of a system including the X-ray detector
and the data acquisition circuitry. The processing circuitry
calculates an X-ray absorption amount of each of a plurality of
base substances based on the count data, concerning the plurality
of energy bands acquired by the data acquisition circuitry, an
energy spectrum of the incident X-rays generated by the X-ray tube,
and the response function read out from the memory circuitry.
[0014] A photon counting apparatus according to the embodiment will
now be described with reference to the accompanying drawings.
[0015] The photon counting apparatus according to the Embodiment is
applicable to any one of an X-Ray CT type apparatus hereinafter)
and an X-ray photography type apparatus to be referred to as a
photon counting XR apparatus hereinafter). The photon counting
apparatus according to the embodiment will be described below in
detail using a photon counting CT apparatus as a detailed
example.
[0016] Various types of photon counting CT apparatuses can be
assumed, including a rotate/rotate-type apparatus in which an X-ray
tube and an X-ray detector integrally rotate around a subject and a
stationary/rotate-type apparatus in which a number of X-ray
detection elements arranged in a ring are fixed, and only an X-ray
tube rotates around a subject. This embodiment is applicable to any
type. However, in the following explanation, the photon counting CT
apparatus is assumed to be a rotate/rotate-type apparatus.
[0017] As the data acquisition method of the photon counting CT
apparatus, a sinogram mode in which the count number of X-ray
photons in each view is counted and a list mode in which the energy
value for each X-ray photon is recorded time-serially are known.
This embodiment is applicable to any type. A photon counting CT
apparatus of the sinogram mode will be exemplified below.
[0018] FIG. 1 is a block diagram showing the arrangement of the
photon counting CT apparatus according to the embodiment. As shown
in FIG. 1, the photon counting CT apparatus according to this
embodiment includes a gantry 10 and a console 30. The gantry 10
supports a rotating frame 11 having a cylindrical shape rotatably
about a rotation axis Z. An X-ray generate on system 13 and an
X-ray detection system 15 are attached to the rotating frame 11 so
as to face each other with respect to the rotation axis Z. A FOV
(Field Of View) is set for the bore of the rotating frame 11. A top
plate 17 is inserted into the bore of the rotating frame 11. A
subject S is placed on the top plate 17. The top plate 17 is
positioned such that the imaging portion of the subject S placed on
the top plate 17 is included in the FOV. The rotating frame 11
receives power from a rotation driver 19 and rotates about the
rotation axis Z at a predetermined angular velocity. The rotation
driver 19 generates the power to rotate the rotating frame 11 in
accordance with a control signal from a gantry control circuitry
21.
[0019] The X-ray generation system 13 generates X-rays in
accordance with a control signal from the gantry control circuitry
21. More specifically, the X-ray generation system 13 includes an
X-ray tube 131 and a high voltage generator 133. Upon receiving
high voltage application and filament current supply from the high
voltage generator 133, the X-ray tube 131 generates X-rays. The
high voltage generator 133 applies a high voltage according to a
control signal from the gantry control circuitry 21 to the X-ray
tube 131, and supplies a filament current according to a control
signal from the gantry control circuitry 21 to the X-ray tube
131.
[0020] The X-ray detection system 15 detects X-rays generated by
the X-ray generation system 13 and transmitted through the subject
S, and acquires, for a plurality of energy bands, count data that
expresses the number of detected X-rays. More specifically, the
X-ray detection system 15 includes an X-ray detector 151 and a data
acquisition circuitry 153.
[0021] The X-ray detector 151 detects X-rays generated by the X-ray
tube 131 and transmitted through the subject S. The X-ray detector
151 includes a plurality of X-ray detection elements that are
two-dimensionally arranged. More specifically, the X-ray detector
151 is assumed to be an indirect type detector. In this case, each
X-ray detection element includes a phosphor (scintillator) that
converts X-rays into fluorescence, and a photodetector that
converts the fluorescence into an electric signal. In this
embodiment, the scintillator detects X-ray photons from the X-ray
tube 131, and generates fluorescent photons in a number
corresponding to the energy of the detected X-ray photons. The
plurality of fluorescent photons are detected by the photodetector.
The photodetector converts the plurality of detected fluorescent
photons into a current signal by photoelectric conversion and
amplifies the current signal. The current signal (electric signal)
from the photodetector is supplied to the data acquisition
circuitry 153. The electric signal has a peak value corresponding
to the energy of the incident X-ray photons. The scintillator
according to this embodiment can contain, as a material any
existing luminescent material, for example, LaBr3 which generates
fluorescence in reaction to X-rays.
[0022] Note that as the X-ray detector 151 according to this
embodiment, not an indirect type detector but a direct type
detector may be used. As the direct type X-ray detector 151, for
example, a type including a semiconductor diode formed by attaching
electrodes to the two ends of a semiconductor is applicable. X-ray
photons that have entered the semiconductor are converted into
electron-hole pairs. The number of electron-hole pairs generated by
entering of one X-ray photon depends on the energy of the incident
X-ray photon. The electrons and holes are attracted by the pair of
electrodes formed at the two ends of the semiconductor. The pair of
electrodes generates electrical pulses having a peak value
corresponding to the charge of an electron-hole pair. One
electrical pulse has a peak value corresponding to the energy of
the incident X-ray photon.
[0023] The data acquisition circuitry 153 acquires count data that
expresses the count number of X-rays detected by the X-ray detector
151 for a plurality of energy bands in accordance with a control
signal from the gantry control circuitry 21. The count data,
concerning the plurality of energy bands corresponds to an energy
spectrum concerning the incident X-rays on the X-ray detector 151,
which is deformed in accordance with the response characteristic of
the X-ray detection system 15. The response characteristic of a
system (standard detection system) including an X-ray detector and
a data acquisition circuitry will be referred to as a detector
response characteristic hereinafter,
[0024] A wedge filter 23 is attached to the X-ray tube 131. The
wedge filter 23 is an X-ray filter used to almost uniform the
spatial dose distribution of X-rays that enter the X-ray detector
151. The wedge filter is formed by a substance having a relatively
small atomic number, for example, aluminum. Typically, the wedge
filter 23 is formed so as to be thick from the center to the ends
in the channel direction of the X-ray detector 151. Note that the
wedge filter 23 may be omitted if unnecessary.
[0025] The gantry control circuitry 21 generally controls various
devices on the gantry 10. For example, the gantry control circuitry
21 controls the X-ray generation system 13, the X-ray detection
system 15, and the rotation driver IS to execute photon counting CT
imaging of the subject S. The rotation driver 19 rotates at a
predetermined angular velocity under the control of the gantry
control circuitry 21. The high voltage generator 133 of the X-ray
generation system 13 applies 4 high voltage corresponding to a set
tube voltage value to the X-ray tube 131 and supplies a filament
current to the X-ray tube 131 under the control of the gantry
control circuitry 21. The data acquisition circuitry 153 of the
X-ray detection system 15 acquires count data on a view basis for
each of a plurality of energy bands in synchronism with view
switching under the control of the gantry control circuitry 21.
[0026] As hardware resources, the gantry control circuitry 21
includes a processor such as a CPU (Central Processing Unit) or MPU
(Micro Processing Unit), and memories such as a ROM (Read Only
Memory) and RAM (Random Access Memory). The gantry control
circuitry 21 may be provided on the gantry 10, the console 30, or a
device separated from the gantry 10 and the console 30. The gantry
control circuitry 21 may be implemented by an application specific
integrated circuitry (ASIC), a field programmable logic device
(FPGA), another complex programmable logic device (CPLD), or a
simple programmable logic device (SPLD), The processor implements
the above-described function by reading out a program saved in the
memories and executing it. Note that instead of saving the program
in the memories, the program may directly be incorporated in a
circuitry of the processor. In this case, the processor implements
the above-described function by reading out the program
incorporated in the circuitry and executing it.
[0027] The console 30 includes a count data memory circuitry 31, a
reconstruction circuitry 33, an I/F circuitry 35, a display
circuitry 37, an input circuitry 39, a main memory circuitry 41,
and a system control circuitry 43, The count data memory circuitry
31, the reconstruction circuitry 33, the I/F circuitry 35, the
display circuitry 37, the input circuitry 33, the main memory
circuitry 41, and the system control circuitry 43 are connected via
a bus.
[0028] The count data memory circuitry 31 is memories such as an
HDD (Hard Disk Drive), an SSD (Solid State Drive), or an integrated
circuitry memory device. More specifically, the count data memory
circuitry 31 stores count data concerning a plurality of energy
bands, which is transmitted from the gantry 10. The count data
memory circuitry 31 may also store data of an X-ray absorption
amount calculated by an X-ray absorption amount calculation module
335. The X-ray absorption amount will be described later.
[0029] The reconstruction circuitry 33 reconstructs a photon
counting CT image concerning the subject S based on the count data.
More specifically, the reconstruction circuitry 33 includes a
response function memory circuitry 331. The response function
memory circuitry 331 scores data of a response function that
associates incident X-rays on the X-ray detector 151 which the
detector response characteristic. The response function defines the
relationship between detection energy for each incident X-ray and
the output response of the system. For example, the response
function defines the relationship between detection intensity and
detection energy for each incident X-ray, The detection energy
corresponds to the energy of X-ray photons measured by the standard
detection system in response to detection of X-ray photons having
the incident X-ray energy. More specifically, the detection energy
is a value obtained by multiplying the peak value of an analog
electric signal input to a discrimination, circuitry (to be
described later) or the data value of a digital signal by a
predetermined conversion factor. The detection, intensity
corresponds to the intensity of X-rays having the incident X-ray
energy, in other words, the count number of X-ray photons. The
response function is generated in advance by a response function
generation module 333 or another processor. Details of the response
function will be described later.
[0030] The standard detection system indicates a system formed from
an X-ray detector and a data acquisition circuitry used to acquire
an actual measured value for response function generation. To
attain a high substance discrimination capability, the standard
detection system and the X-ray detection, system 15 according to
this embodiment preferably have the same structure. For example,
all factors that affect the detector response characteristic, such
as the scintillator material, circuitry arrangement, and sampling
speed are preferably identical in the standard detection system and
the X-ray detection system 15 according to this embodiment. Note
that some of these factors may be different if the influence on the
detector response characteristic is small. The X-ray detection
system 15 included in the photon counting CT apparatus according to
this embodiment may be used to acquire an actual measured value for
response function generation. In this case, the standard detection
system indicates the X-ray detection system 15.
[0031] As shown in FIG. 1, the reconstruction circuitry 33 may
include, as hardware resources, a processor such as a CPU, MPU, or
GPU (Graphics Processing Unit), and memories such as a ROM and RAM,
in addition to the response function memory circuitry 331. The
processing unit implements a response function generation module
333, an X-ray absorption amount calculation module 335, and a
reconstruction processing module 337 by reading out a
reconstruction program saved in the memories and executing it.
[0032] By executing the response function generation module 333,
the reconstruction circuitry 33 generates data of a response
function that expresses the detector response characteristic. For
example, the reconstruction circuitry 33 measures the response
(that is, detection energy and detection intensity) of the standard
detection system to a plurality of monochromatic X-rays having a
plurality of incident X-ray energies by predictive calculations,
experiments, and a combination of predictive calculations and
experiments, and generates a response function based on the
measured values of detection energy and detection intensity. The
reconstruction circuitry 33 may generate data of a response
function based on actual measured values acquired in calibration or
the like. The generated data of the response function is stored in
the response function memory circuitry 331. Note that the data of
the response function to be stored in the response function memory
circuitry 331 need not always be generated by the reconstruction
circuitry 33. The data of the response function may be generated by
a computer apparatus in another facility. In this case, the data
may be transmitted from the computer apparatus to the photon
counting CT apparatus according to this embodiment, or read out
from a portable memory medium that stores the data of the response
function to the photon counting CT apparatus according to this
embodiment.
[0033] By executing the X-ray absorption amount calculation module
335, the reconstruction circuitry 33 calculates an X-ray absorption
amount concerning each of a plurality of base substances based on
count data concerning the plurality of energy bands, the energy
spectrum of the incident X-rays on the subject S, and the response
function stored in the response function memory circuitry 331. The
reconstruction circuitry 33 calculates the X-ray absorption amount
based on the count data and the energy spectrum of the incident
X-rays on the subject S using the response function, thereby
calculating an X-ray absorption amount without the influence of the
response characteristic of the X-ray detection system 15.
Processing of obtaining the X-ray absorption amount for each base
substance is also called substance discrimination. All substances
such as calcium, calcification, bone, fat, muscle, air, organ,
lesion, hard tissue, soft tissue, and contrast substance can be set
as the base substance. The type of the calculation target base
substance is decided in advance by a user or the like via the input
circuitry 39 and the like. The X-ray absorption amount represents
the amount of X-rays absorbed by the base substance. More
specifically, the X-ray absorption amount is defined by a
combination of an X-ray attenuation coefficient and an X-ray
transmission path length.
[0034] By executing the reconstruction processing module 337, the
reconstruction circuitry 33 reconstructs a photon counting CT image
that expresses the spatial distribution of an imaging target base
substance out of a plurality of base substances, based on the X-ray
absorption amount concerning each of the plurality of base
substances calculated by the X-ray absorption amount calculation
module 335. The imaging target base substance can include one type
of base substance or a plurality of types of base substances. The
imaging target, base substance can be set via the input circuitry
39 or automatically arbitrarily.
[0035] Note that the response function generation module 333, the
X-ray absorption amount calculation module 335, and the
reconstruction processing module 33 are assumed to be implemented
by executing the reconstruction program by the processing unit.
However, the embodiment is not limited to this, for example, the
reconstruction circuitry 33 may include a processing circuitry for
the response function generation module 333, a processing circuitry
for the X-ray absorption amount calculation module 335, and a
processing circuitry for the reconstruction processing module 337,
Each of these processing circuitry may be implemented, by ASIC,
FPGA, CPLD, or SPLD.
[0036] The I/F circuitry 35 is an interface for communication
between the console 30 and the gantry 10. For example, the I/F
circuitry 35 supplies an imaging start signal, imaging stop signal,
and the like from the system control circuitry 43.
[0037] The display circuitry 3 displays a photon counting CT image
or the like on a display device. As the display device, for
example, a CRT display, a liquid crystal display, an organic EL
display, or a plasma display can appropriately be used.
[0038] The input circuitry 39 receives various kinds of
instructions and information input, from the user via an input
device. As the input device, a keyboard, a mouse, various kinds of
switches, and the like are usable.
[0039] The main memory circuitry 41 is a memory device that stores
various kinds of information. For example, the main memory
circuitry 41 stores the image generation program of a photon
counting CT image according to this embodiment, and the like.
[0040] The system control circuitry 43 functions as the center of
the photon counting CT apparatus according to this embodiment. The
system control circuitry 43 reads out an imaging program, according
to this embodiment from the main memory circuitry 41, and controls
various kinds of constituent elements in accordance with the
imaging program. Photon counting CT imaging for generating a photon
counting CT image according to this embodiment is thus
performed.
[0041] The response function according to this embodiment will be
described next.
[0042] Generally, in dual energy CT, substance discrimination is
performed in accordance with below equation (1).
I.sub.det(E)=I.sub.0(E)exp(-.mu..sub.0(E)L.sub.0-.mu..sub.1(E)L.sub.1)
(1)
[0043] In the equation (1), E is the energy of X-rays, and
I.sub.det(E) is the energy spectrum of X-rays measured by the
standard detection system. Mote that, the energy spectrum indicates
the energy distribution of X-ray intensities. I.sub.det(E) is the
energy spectrum of X-rays that enter the subject S. When the wedge
filter is used, I.sub.0(E) represents the energy spectrum of X-rays
that enter the object after passing through the wedge filter. When
the wedge filter is not used, I.sub.0(E) represents the energy
spectrum of X-rays emitted by the X-ray tube. In addition,
.mu..sub.0(E) is the X-ray attenuation coefficient of a base
substance 0, and L.sub.0 is the path length (transmission path
length) of X-rays transmitted through the base substance 0.
Similarly, .mu..sub.1(E) is the X-ray attenuation coefficient of a
base substance 1, and L.sub.1 is the path length (transmission path
length) of X-rays transmitted through the base substance 1.
.mu..sub.0(E)L.sub.0 is the X-ray absorption amount concerning the
base substance 0, and .mu..sub.1(E)L.sub.1 is the X-ray absorption
amount concerning the base substance 1. Equation (1) includes two
unknowns .mu..sub.0(E)L.sub.0 and .mu..sub.1(E)L.sub.1 which can be
solved using two equations in theory. In current CT, different two
data sets concerning two different tube voltages are acquired,
thereby obtaining the solution to equation (1). Note that the data
set indicates a set of data representing (E) and I.sub.0(E).
[0044] In photon counting CT as well, substance discrimination can
be performed based on the same concept as the dual energy CT. In
photon counting CT, one set of count data is used because data
acquisition is performed using one tube voltage. However, when the
energy band is divided into two parts, substance discrimination can
be done based on two count data sets concerning the two energy
bands. Independence of the two count data sets is a precondition
for the concept. A direct type detector having an excellent energy
resolving power can meet this precondition. However, an indirect
type detector having a low energy resolving power cannot meet the
precondition in many cases because of low independence.
[0045] Equation (1) actually holds for an ideal X-ray detection
system. In actuality, substance discrimination can be described as
below equation (2).
I.sub.det(E)={I.sub.0(E)exp(-.mu..sub.0(E)L.sub.0-.mu..sub.1(E)L.sub.1)}-
R(E) (2)
[0046] In the equation (2), is a convolution operator. The
right-hand side is called a model function. The model function
describes the detector response characteristic of the X-ray
detection system 15, R(E) is a response function representing the
detector response characteristic. The expression in braces
represents the energy spectrum of X-rays immediately before
entering the X-ray detection system 15. Equation (2) convolutes the
energy spectrum of X-rays immediately before entering the X-ray
detection system 15 by the response function of the X-ray detection
system 15, thereby obtaining the energy spectrum of the X-rays
measured by the X-ray detection system 15.
[0047] Here, the response function describes the detector response
characteristic of the standard detection system to monochromatic
X-rays. An ideal X-ray detector generates an output signal like a
delta function having only a detection energy corresponding to
incident X-ray energy. However, the output signal of an actual
X-ray detector is distributed not as a delta function but as a
Gaussian function because of the energy resolving power. In
addition, the output signal often exhibits a complex structure such
as a deviation from the Gaussian function or a continuous component
caused by components on the low energy side (G. F. Knoll,
"Radiation Detection and Measurement, Third Edition", Nikkan Kogyo
Shimbun, 2001). Under a high dose, the response function further
deforms due to pile-up, polarization, or the like.
[0048] As described above, since the indirect type detector has a
low energy resolving power, the low independence between adjacent
energy bands in equation (1) poses a problem. However, when the
response function is correctly considered like equation (2),
.mu..sub.0(E)L.sub.0 and .mu..sub.1(E)L.sub.1 can accurately be
estimated.
[0049] A problem here is the number of energy bands. According to
the concept of equation (1), energy bands as many as the unknowns
suffice, and at least two energy bands suffice. In equation (2),
however, many energy bands are needed to consider the response
function.
[0050] FIG. 2 is a graph showing the energy spectrum I.sub.0(E) of
X-rays generated by the X-ray tube without the wedge filter. In the
graph of FIG. 2, the ordinate is defined as the count number, and
the abscissa is defined as the energy [keV]. The anode target
substance is tungsten, and the tube voltage is 120 kV. As shown in
FIG. 2, the energy spectrum of the X-rays generated by the X-ray
tube includes characteristic X-rays resulting from the anode target
substance as well as continuous X-ray components from 0 keV to 120
keV corresponding to the tube voltage.
[0051] In actual photon counting CT imaging, the subject S exists
between the X-ray tube and the X-ray detector. For this reason, the
X-rays having the energy spectrum shown in FIG. 2 are attenuated by
absorption to the subject 8 and then detected by the X-ray
detection system 15. The energy spectrum of the detected X-rays is
convoluted by the response function corresponding to the detector
response characteristic.
[0052] FIG. 3 shows the output of the standard detection system
using LaBr3 as a scintillator with respect to monochromatic X-rays.
In FIG. 3, the ordinate is defined as the intensity [A.U.], and the
abscissa is defined as the energy [keV]. Out of LaBr3, lanthanum La
mainly reacts to X-rays. The k-edge of lanthanum is 38.9 keV. FIG.
3 shows the output of the standard detection system with respect to
monochromatic X-rays of 30 keV. In this case, since the energy (30
keV) of the monochromatic X-rays that have entered the standard
detection system is lower than the energy (38.9 keV) of the k-edge
of lanthanum, the output of the standard detection system is
distributed with a peak at the energy corresponding to 30 keV. The
peak corresponding to the energy of the incident monochromatic
X-rays is called a main peak. The energy width of the main peak
represents the energy resolving power of the standard detection
system concerning the energy of the main peak.
[0053] FIG. 4 shows the output of the standard detection system
using LaBr3 as a scintillator with respect to other monochromatic
X-rays. FIG. 4 shows the output of the standard detection system
with respect to monochromatic X-rays of 50 keV. In this case, since
the energy (50 keV) of the monochromatic X-rays that have entered
the standard detection system is higher than the energy (38.9 keV)
of the k-edge of lanthanum, the output of the standard detection
system is distributed not only with the main peak at the energy
corresponding to 50 keV but also with a peak near 2.7 keV. This
peak is called an escape peak. When the energy of monochromatic
X-rays that have entered the standard defection system is higher
than the energy of the k-edge of lanthanum, the lanthanum which
absorbs (photoelectrically absorbs) the X-rays and jumps an excited
state often emits the k-characteristic X-rays of lanthanum when the
lanthanum fails from the excited state to the ground state. If the
k-characteristic X-rays escape without being absorbed by the
scintillator, the scintillator absorbs an energy (about 17 keV)
obtained by subtracting the energy (about 33 keV) of the
k-characteristic X-rays from the incident X-ray energy. The escape
peak is a peak corresponding to the energy (about 17 keV)
equivalent to the difference between the incident X-ray energy and
the energy (about 33 keV) of the k-characteristic X-rays.
[0054] FIG. 5 is a graph schematically showing the response
function according to this embodiment. The response function
according to this embodiment is defined by, for example, a function
including the incident X-ray energy [keV]f the detection energy
[keV], and the detection intensity [A.U.] as variables. In the
graph of FIG. 5, the detection intensity is assigned to an
orthogonal coordinate system in which the ordinate (y-axis) is
defined as the incident X-ray energy, and the abscissa (x-axis) is
defined as the energy [keV]. In FIG. 5, the detection intensity is
expressed by shading.
[0055] In other words, the response function according to this
embodiment defines the relationship between the detection energy
and the detection intensity for each of a plurality of incident
X-ray energies. For example, a response function representing the
detector response characteristic for a certain incident X-ray
energy is expressed as a cross section of the graph of FIG. 5 at
the incident X-ray energy. As shown in FIG. 5, the main peak is
observed at a point corresponding to Y=X. The dotted line
corresponds to the energy of the k-edge of lanthanum. In the
incident X-ray energies higher than the energy of the k-edge of
lanthanum, escape peaks are observed at a predetermined interval on
the low detection energy side of the main peak. The detection
energy difference between the main peak and the escape peak
corresponds to the detection energy equivalent to the energy of the
k-characteristic X-rays of lanthanum. In addition, the peak of
k-characteristic X-rays when the k-characteristic X-rays of
lanthanum generated by another scintillator are absorbed by the
scintillator is observed. Since the energy of the k-characteristic
X-rays is constant, almost the same detection energy, that is, a
detection energy parallel to the y-axis is observed independently
of the incident X-ray energy. As described above, the response
function according to this embodiment expresses a main peak
corresponding to the incident X-ray energy, an escape peak
corresponding to the incident X-ray energy to the energy of the
k-characteristic X-rays of the scintillator material, and a peak
corresponding to the energy of the k-characteristic X-rays of the
scintillator material.
[0056] In the response function generation module 333 according to
this embodiment, the reconstruction circuitry 33 measures the
response of the standard detection system to a plurality of
monochromatic X-rays having a plurality of incident X-ray energies
by predictive calculations, experiments, and a combination of
predictive calculations and experiments, and generates a response
function based on the measured data. More specifically, the X-ray
source irradiates the X-ray detector with monochromatic X-rays, and
the X-ray detector detects the monochromatic X-rays. The detection
energy and detection intensity of the X-ray detector are measured
by an existing measurement device. The monochromatic X-rays are
emitted sequentially from the lower limit energy to the upper limit
energy of the energy range necessary for the response function. In
the response function generation module 333, the reconstruction
circuitry 33 records the measured detection energy and detection
intensity for each incident X-ray energy of the emitted
monochromatic X-rays. Note that if the wedge filter is used in the
photon counting CT imaging, the standard detection system
preferably detects the monochromatic X-rays transmitted through the
wedge filter. In this case, in the response function generation
module 333, the reconstruction circuitry 33 records the detection
energy and detection intensity of the X-ray detector tor the
monochromatic X-rays transmitted through the wedge filter for each
incident X-ray, The record of the detection energy and detection
intensity for each incident X-ray is generated as the response
function. Note that as for the response function, a mathematical
model of a response function may be formed by predictive
calculation, and the mathematical model may be corrected based on
experimental values in a facility where high-intensity
monochromatic X-rays such as synchrotron radiation can be
obtained.
[0057] The energy spectrum of X-rays output from the X-ray detector
has a shape obtained by attenuating the energy spectrum shown in
FIG. 2 by object absorption in accordance with equation (2) and
convoluting the energy spectrum of the X-rays that have undergone
the object absorption by the response function shown in FIG. 1. The
energy spectrum of the X-rays that have undergone the object
absorption is expressed by the model function of equation (2).
[0058] FIG. 6 is a graph showing comparison between an energy
spectrum represented by count data of actual measurement and an
energy spectrum represented by the model function. The ordinate of
the upper section of FIG. 6 is defined as the detection intensity
(count number), and the abscissa represents the energy. The
ordinate and abscissa are expressed as logarithms. Referring to
FIG. 6, the range from 0 keV to 200 keV is divided into 128 energy
bands. The widths of the energy bands are set identically. In the
graph of the upper section, each count data of actual measurement
is indicated, by a cross, and the model function is indicted by a
solid line. The lower section is a graph representing the
difference (residual) between the energy spectrum represented by
the count data of actual measurement and the energy spectrum
represented by the model function. Data acquisition was conducted
using a tube voltage of 120 kV and without a wedge filter and a
subject (air). As can be seen from FIG. 6, the model function can
generally reproduce the count data of actual measurement. In
actuality, since a subject exists, an absorption structure further
appears. The subject absorption can also be incorporated as a model
function by assuming several base substances in accordance with
equation (2).
[0059] However, as is apparent from FIG. 6, in a case where a base
substance is quantitatively evaluated by convoluting a response
function, if data sets, that is, energy bands as many as the
unknowns are simply provided, a plurality of sets of solutions that
meet equation (2) are obtained, and the substance and the
absorption amount cannot be identified. To obtain a solution in
accordance with equation (2), at least 16, if possible, 50 or more
energy bands preferably exist, as shown in FIG. 6.
[0060] This analysis scheme can cope with the below equation (3) in
which the number of base substances is extended to three from
equation (2).
I.sub.det(E)
={I.sub.0(E)exp(-.mu..sub.0(E)L.sub.0-.mu..sub.1(E)L.sub.1-.mu..sub.2(E)L-
.sub.2)}R(E) (3)
[0061] In the conventional dual energy CT, the k-edge of a base
substance cannot be taken into consideration. However, in the image
reconstruction method according to this embodiment, a k-edge
structure can easily be reproduced by including the k-edge in the
model function. Hence, by actively using a substance having a
k-edge within the detection energy range, the image reconstruction
method according to this embodiment can be applied to a k-edge
imaging method as well, which enables high-contrast imaging that
cannot be implemented by the dual energy CT.
[0062] The structure of the data acquisition circuitry 153 capable
of setting many energy bands will be described below.
[0063] FIG. 7 is a block diagram showing an example of the
arrangement of a data acquisition circuitry 153-1 according to this
embodiment. Note that the data acquisition circuitry 153-1 includes
readout channels as many as channels corresponding to the X-ray
detection elements. The plurality of readout channels are
parallelly implemented on an integrated circuitry such as an ASIC.
FIG. 7 illustrates only the arrangement of the data acquisition
circuitry 153-1 corresponding to one readout channel to avoid
redundancy.
[0064] The data acquisition circuitry 153-1 shown in FIG. 7
includes a preamplifier circuitry 61, a waveform shaping circuitry
63, a plurality of pulse-height discrimination circuitry 65, a
plurality of counting circuitry 67, and an output circuitry 69.
[0065] The preamplifier circuitry 61 amplifies a current, signal
from the X-ray detection element of a connection destination. More
specifically, the preamplifier circuitry 61 converts a current
signal from the X-ray detection element of the connection
destination into a voltage signal, having a voltage value (peak
value) proportional to the charge amount of the current signal. The
waveform shaping circuitry 63 is connected to the preamplifier
circuitry 61. The waveform shaping circuitry 63 shapes the waveform
of the voltage signal from the preamplifier circuitry 61. More
specifically, the waveform shaping circuitry 63 reduces the pulse
width of the voltage signal from the preamplifier circuitry 61.
[0066] A plurality of counting channels corresponding to the number
of energy bands are connected to the waveform shaping circuitry 63.
When n energy bands are set, n counting channels are provided. More
specifically, n is preferably 16 or more, as described above. Each
counting channel includes a pulse-height discrimination circuitry
65-n and a counting circuitry 61-n.
[0067] Each pulse-height discrimination circuitry 65-n
discriminates the peak value of the voltage signal from the
waveform shaping circuitry 63, that is, the energy of X-ray photons
detected by the X-ray detection element. More specifically, the
pulse-height discrimination circuitry 65-n includes a D/A
conversion circuitry (DAC) 651-n and a comparison circuitry 653-n.
The DAC 651-n inputs a digital signal (to be referred to as a
digital threshold signal hereinafter) having a data value
corresponding to an energy threshold from the gantry control
circuitry 21 (not shown). The DAC 651-n converts the input digital
threshold signal into an analog electric signal (to be referred to
as an analog threshold signal hereinafter) having a peak value
corresponding to the data value (energy threshold) of the digital
threshold signal. Digital threshold signals corresponding to
different thresholds are supplied from, for example, the gantry
control circuitry 21 to the DACs 651-n. If the voltage signal from
the waveform shaping circuitry 63 has a peak value corresponding to
an energy band corresponding to the peak value (energy threshold)
of the analog threshold signal from the DAC 651-n, the comparison
circuitry 653-n outputs an electrical pulse signal. For example, if
the peak value of the electrical pulse from the waveform shaping
circuitry 63 is the peak value corresponding to an energy band
bin1, a comparison circuitry 653-1 for the energy band bin1 outputs
an electrical pulse signal. On the other hand, if the peak value of
the electrical pulse from the waveform shaping circuitry 63 is not
the peak value corresponding to the energy band bin1, the
comparison circuitry 653-1 for the energy band bin1 does not output
an electrical pulse signal.
[0068] The counting circuitry 67-n counts the electrical pulse
signal from the pulse-height discrimination circuitry 65-n at a
readout period that matches the view switching period. More
specifically, the gantry control circuitry 21 supplies a trigger
signal to the counting circuitry 67-n at the switching timing of
each view. Along with the supply of the trigger signal, the
counting circuitry 67-n adds 1 to the count number stored in the
internal memory every time an electrical pulse signal is input from
the pulse-height discrimination circuitry 65-n. Along with the
supply of the next trigger signal, the counting circuitry 67-n
reads out the data of the count number (that is, count data)
accumulated in the internal memory, and supplies it to the output
circuitry 69. The counting circuitry 67-n also sets again the count
number accumulated in the internal memory to the initial value
every time a trigger signal is supplied. The counting circuitry
67-n thus counts the electrical pulse signal on a view basis.
[0069] The output circuitry 69 is connected to the counting
circuitry 67-n as many as the plurality of readout channels
included in the X-ray detector 151. For each of the plurality of
energy bands, the output circuitry 65 integrates the count data
from the counting circuitry 67-n as many as the plurality of
readout channels, and generates count data for the plurality of
readout channels on a view basis. The count data of each energy
band is a set of data of count numbers defined by a channel, a
segment (column), and an energy band. The count data of each energy
band is transmitted to the console 30 on a view basis. The count
data on a view basis is called a count data set.
[0070] As described above, the data acquisition circuitry 153-1
shown in FIG. 7 includes the analog pulse-height discrimination
circuitry 65-n and the counting circuitry 67-n, which are the same
as the components of a conventional data acquisition circuitry, in
the data acquisition circuitry 153-1 shown in FIG. 7, the number of
counting channels is extended as compared to the conventional data
acquisition circuitry, However, the individual components have been
field-proven. Although the implementation area and power increase
in proportion to the number of counting channels, the circuitry can
be implemented.
[0071] FIG. 8 is a block diagram showing another example of the
arrangement of a data acquisition circuitry 153-2 according to this
embodiment. Note that the data acquisition circuitry 153-2 shown in
FIG. 8 also includes readout channels as many as channels
corresponding to the X-ray detection elements, like the data
acquisition circuitry 153-1 shown in FIG. 7. The plurality of
readout channels are parallelly implemented on an integrated
circuitry such as an ASIC. FIG. 8 illustrates only the arrangement
of the data acquisition circuitry 153-2 corresponding to one
readout channel to avoid redundancy.
[0072] The data acquisition circuitry 153-2 shown in FIG. 8
includes a preamplifier circuitry 71, a variable gain amplifier
circuitry 73, a buffer amplifier circuitry 75, an A/D conversion
circuitry (to be referred to as an ADC hereinafter) 77, a digital
filter circuitry 79, an integrated counting circuitry 81, and an
output circuitry 83.
[0073] The preamplifier circuitry 71 amplifies a current signal
from the X-ray detection element of a connection destination. More
specifically, the preamplifier circuitry 71 converts a current
signal from the X-ray detection element of the connection
destination into a voltage signal having a voltage value (peak
value) proportional to the charge amount of the current signal. The
variable gain amplifier circuitry 73 is connected to the
preamplifier circuitry 71. The variable gain amplifier circuitry 73
amplifies the voltage signal from the preamplifier circuitry 71 by
a variable gain. The gain of the variable gain amplifier circuitry
73 can be set to an arbitrary value by, for example, the user via
the input circuitry 39. The buffer amplifier circuitry 75 is
connected to the variable gain amplifier circuitry 73. The buffer
amplifier circuitry 75 amplifies the voltage signal from the
variable gain amplifier circuitry 73 by a gain to suppress a
variation in the frequency in the ADC 77 of the subsequent stags.
The ADC 77 is connected to the buffer amplifier circuitry 75.
[0074] The ADC 77 samples the voltage signal from the buffer
amplifier circuitry 75 by a predetermined number of bits, and
converts the voltage signal into a discrete time series digital
signal having a data value corresponding to the peak value of the
voltage signal from the buffer amplifier circuitry 75. The digital
filter circuitry 79 is connected to the ADC 77. The digital filter
circuitry 79 analyzes the digital signal from the ADC 77, thereby
specifying the arrival time of X-ray photons and the energy of the
X-ray photons. The arrival time of the X-ray photons corresponds to
the time at which the peak is recorded, and the energy of the X-ray
photons corresponds to the data value at the peak. A digital signal
representing the arrival time of X-ray photons and the energy of
the X-ray photons will be referred to as an energy signal
hereinafter. The integrated counting circuitry 81 is connected to
the digital filter circuitry 79.
[0075] The integrated counting circuitry 81 includes counting
channels in a number matching the number n of energy bands. Each
counting channel includes a discrimination circuitry 811-n and a
counting circuitry 813-n. Based on an energy signal repetitively
supplied from the digital filter circuitry 79, the plurality of
discrimination circuitry 811-n discriminate the energy band to
which X-ray photons corresponding to the energy signal belong.
Different energy thresholds are assigned to the discrimination
circuitry 811-n. Each discrimination circuitry 811-n performs
threshold processing based on the energy threshold for a
repetitively supplied energy signal so as to pass an energy signal
belonging to an energy band corresponding to the energy threshold
and block an energy signal belonging to an energy band that does
not correspond to the energy threshold. Each counting circuitry
813-n counts the energy signal supplied from the discrimination
circuitry 811-n of the connection source. The plurality of counting
circuitry 813-n count the energy signals from the discrimination
circuitry 811-n at a readout period that matches the view switching
period. More specifically, the gantry control circuitry 21 supplies
a trigger signal to the counting circuitry 813-n at the switching
timing of each view. Along with the supply of the trigger signal,
the counting circuitry 813-n adds 1 to the count number stored in
the internal memory every time an energy signal is input from the
discrimination circuitry 811-n. Along with the supply of the next
trigger signal, the counting circuitry 813-n reads out the data of
the count number (that is, count data) accumulated in the internal
memory, and supplies it to the output circuitry 83. The counting
circuitry 813-n also sets again the count number accumulated in the
internal memory to the initial value every time a trigger signal is
supplied. The counting circuitry 813-n thus counts the energy
signal on a view basis.
[0076] The output circuitry 83 is connected to the integrated
counting circuitry 81 corresponding to the plurality of readout
channels included in the X-ray detector 151. for each of the
plurality of energy bands, the output circuitry 83 integrates the
count data from the integrated counting circuitry 81 corresponding
to the plurality of readout channels, and generates count data for
the plurality of readout channels on a view basis. The count data
of each energy band is a set of data of count numbers defined by a
channel, a segment (column), and an energy band. The count data of
each energy band is transmitted to the console 30 on a view basis.
The count data on a view basis is called a count data set.
[0077] As described above, in the data acquisition circuitry 153-2
shown in FIG. 8, the ADC 7 is implemented in the ASIC, the output
from the ADC 77 is processed by the digital filter circuitry 79,
and the output from the digital filter circuitry 79 is output as
count data representing an energy spectrum. The X-ray intensity
required in CT is as high as 3-600.times.10.sup.6 ph/s/mm.sup.2
after passing through the wedge filter ("Enabling Photon Counting
Clinical X-ray CT", K. Taguchi, et al. 2009 IEEE Proc. Nucl. Sci.
pp. 3581-3585). In the data acquisition circuitry 153-2 shown in
FIG. 8, the sampling speed of the ADC 77 is required to be 100 Msps
(Sampling Per Second) or more. Hence, there is a fear of heat
generation, implementation area, and the like, as in the
arrangement shown in FIG. 7. However, an ADC with a speed of 1 Gsps
or more is also commercially available, which has a high degree of
implementability.
[0078] An example of the operation of photon counting CT imaging
according to this embodiment will be described next. FIG. 9 is a
flowchart showing the typical procedure of photon counting CT
imaging performed under the control of the system control circuitry
43.
[0079] As shown in FIG. 9, the system control circuitry 43 controls
the gantry control circuitry 21 to perform photon counting CT
imaging for the subject S and acquire count data concerning a
plurality of energy bands (step S1). In step SI, the gantry control
circuitry 21 controls the X-ray generation system 13, the X-ray
detection system 15, and the rotation driver 19 to execute photon
counting CT imaging of the subject S. The rotation driver 19
rotates at a predetermined angular velocity under the control of
the gantry control circuitry 21. The high voltage generator 133 of
the X-ray generation system 13 applies a high voltage corresponding
to a set tube voltage value to the X-ray tube 131 and supplies a
filament current to the X-ray tube 131 under the control of the
gantry control circuitry 21. The data acquisition circuitry 153 of
the X-ray detection system 15 acquires a count data set on a view
basis in synchronism with view switching under the control of the
gantry control circuitry 21. The count data set is transmitted from
the gantry to the console by a transmission apparatus (not shown).
The count data set is a set of data in which the energy value of an
energy band, a segment number, and a channel number are assigned to
each of the plurality of X-ray detection elements. In other words,
the count data set represents an energy distribution for the count
number of each X-ray detection element.
[0080] When step S1 is performed, the system control circuitry 43
causes the reconstruction circuitry 3 to execute the X-ray
absorption amount calculation module 335 (step S2). In step S3, the
reconstruction circuitry 33 calculates the X-ray absorption amount
of each of a plurality of base substances without the influence of
a detector response characteristic based on the count data set
acquired in step S1, the energy spectrum of incident X-rays on the
subject S, and a response function stored in the response function
memory circuitry 331. The X-ray absorption amount calculation
processing for each base substance is also called substance
discrimination.
[0081] More specifically first, the reconstruction circuitry 33
reads out data of the response function from the response function
memory circuitry 331. Next, the reconstruction circuitry 33
calculates the difference between the count data set and a model
function while changing the X-ray absorption amount of each base
substance included in the model, function, and decides the final
X-ray absorption amount of each base substance with which the
calculated difference becomes smaller than a threshold. At this
time, the reconstruction circuitry 33 decides the final X-ray
absorption, amount with which the difference becomes smaller than
the threshold simultaneously for all of the plurality of energy
bands. As described above, the model function is defined by
convolution of the response function for an integrated value of the
energy spectrum of incident X-rays on the subject S and a power of
a Napier's constant e using the X-ray absorption amount as an
exponent. The energy spectrum of the incident X-rays on the subject
S is measured by the X-ray detection system 15 at the time of, for
example, calibration, and stored in the response function memory
circuitry 331, the main memory circuitry 41, or the like. A
threshold is decided to an arbitrary value in advance. For example,
if there are two types of base substances, that is, the base
substances 0 and the base substance 1, the reconstruction circuitry
33 decides the final X-ray absorption amount on a view basis in
accordance with the following procedure.
[0082] First, an initial value .mu..sub.01 of an X-ray attenuation
coefficient and an initial value L.sub.01 of a transmission path
length for the base substance 0 and an initial value .mu..sub.11 of
an X-ray attenuation coefficient and an initial value L.sub.11 of a
transmission path length for the base substance 1 are set
automatically or according to an instruction input by the user via
the input circuitry 35. The initial value of the X-ray attenuation
coefficient and the initial value L.sub.01 of the transmission path
length and the initial value .mu..sub.11 of the X-ray attenuation
coefficient and the initial value L.sub.11 of the transmission path
length are set in advance based on experimental or statistical
knowledge. As indicated by expressions (4) below, the
reconstruction circuitry 33 calculates the difference (left-hand
side of expression (4)) between the count data set I.sub.det(E) and
an initial model function M.sub.0(E), and compares the difference
with the threshold .
.SIGMA..sub.bini.sup.j{I.sub.det(E)-M.sub.0(E).ltoreq..di-elect
cons.}M.sub.0(E)={I.sub.0(E)exp(-.mu..sub.0(E)L.sub.01-.mu..sub.11(E)L.su-
b.11)}R(E) (4)
[0083] As indicated by expressions (4), the reconstruction
circuitry 33 performs the comparison of the difference and the
threshold for all energy bands i. Upon determining for ail of the
plurality of energy bands that the difference is smaller than the
threshold , the initial value .mu..sub.01 of the X-ray attenuation
coefficient and the initial value L.sub.01 of the transmission path
length for the base substance 0 and the initial value .mu..sub.11
of the X-ray attenuation coefficient and the initial value L.sub.11
of the transmission path length for the base substance 1 are
decided as final values. The final X-ray attenuation coefficient
and transmission path length are stored in the count data memory
circuitry 31 for each base substance.
[0084] On the other hand, upon determining for at least one of the
plurality of energy bands that the difference is larger than the
threshold , the X-ray attenuation coefficient and the transmission
path length for the base substance 0 and the X-ray attenuation
coefficient and the transmission path length for the base substance
1 are changed. The X-ray attenuation coefficients and the
transmission path lengths are changed by an existing method. As
indicated by expressions (5) below, the difference (left-hand side
of expression (5)) between the count data set and a model function
M.sub.0(E) after the nth iteration is calculated and compared with
the threshold .
.SIGMA..sub.bini.sup.j{I.sub.det(E)-M.sub.n(E).ltoreq..di-elect
cons.}M.sub.n(E)={I.sub.0(E)exp(-.mu..sub.0n(E)L.sub.0n-.mu..sub.1n(E)L.s-
ub.11)}R(E) (5)
[0085] In the above equations n indicates the nth iteration. Note
that the initial value is n=0.
[0086] In this way, the difference (left-hand side of expression
(5)) between the count data set and the model function M.sub.n(E)
is iteratively calculated until the difference between the count
data set and the model function M.sub.n(E) falls below the
threshold for all of the plurality of energy bands.
[0087] Upon determining for all of the plurality of energy bands
that the difference between the count data set and the model
function M.sub.n(E) fails below the threshold , an X-ray
attenuation coefficient .mu..sub.0n and a transmission path length
L.sub.0n for the base substance 0 and an X-ray attenuation
coefficient .mu..sub.1n and a transmission path length L.sub.1n for
the base substance 1 included in the model function M.sub.n(E) are
decided as final values. The final X-ray attenuation coefficient
and transmission path length are stored in the count data memory
circuitry 31 for each base substance.
[0088] Note that the above-described method of calculating the
X-ray attenuation coefficient and transmission path length for a
base substance is merely an example, and the X-ray attenuation
coefficient and transmission path length for a base substance can
be calculated by any calculation method. For example, the
difference between inverse convolution of the count data set by the
response function and the integrated value of the energy spectrum
of incident X-ray energy and a power of the Napier's constant using
the X-ray absorption amount as an exponent may be compared with the
threshold.
[0089] In the above-described method, the threshold e is set in
advance, and the model function M.sub.n(E) when the difference
between the count data set and the model function M.sub.n(E) falls
below the threshold e is decided as the final model function.
However, the embodiment is not limited to this. For example, the
difference between the count data set and the model function
M.sub.0(E) is calculated a plurality of times, that is, m times,
and out of m model functions M.sub.n(E), a model function M(E) when
the difference is a local minimum is decided as the final model
function.
[0090] In addition, the X-ray absorption amount with which the
difference between the count data set and the model function
becomes smaller than the threshold simultaneously for all of the
plurality of energy bands is decided as the final X-ray absorption
amount. However, the embodiment is not limited to this. For
example, an X-ray absorption amount with which the difference
becomes smaller than the threshold for a predetermined number of
energy bands out of the plurality of energy bands may be decided as
the final X-ray absorption amount.
[0091] When step S2 is performed, the system control circuitry 43
causes the reconstruction circuitry 33 to perform reconstruction
processing (step S3). In step S3, based on the X-ray absorption
amount of each of the plurality of base substances calculated in
step S2, the reconstruction circuitry 33 reconstructs a photon
counting CT image that expresses the spatial distribution of the
base substance included in the subject S. The type of the base
substance to be used for the reconstruction can arbitrarily be
selected from the plurality of base substances whose X-ray
absorption amounts are calculated in step 32. For example, a photon
counting CT image reconstructed based on the X-ray absorption
amount of the base substance 0 expresses the spatial distribution
of the base substance 0. A photon counting CT image reconstructed
based on the X-ray absorption amount of the base substance 1
expresses the spatial distribution of the base substance 1. Note
that as the image reconstruction algorithm, an existing image
reconstruction algorithm, such as an analytic image reconstruction
method based on FBP (Filtered Back Projection) or CBP (Convolution
Back Projection) or a statistical image reconstruction method based
on RL-EM (Maximum Likelihood Expectation Maximization) OS-EM
(Ordered Subset Expectation Maximization), or OS-SART (Ordered
Subset. Simultaneous Algebraic Reconstruction Techniques) is used.
A k-edge imaging method may be incorporated in the image
reconstruction algorithm. Plainly speaking, k-edge imaging is a
method of reconstructing a photon counting CT image expressing the
spatial distribution of an imaging target substance based on count
data concerning energy bands on both sides of an energy band to
which the k-edge or the imaging target substance belongs.
[0092] When step S3 is performed, the system control circuitry 43
causes the display circuitry 37 to perform display processing (step
S4). In step S4, the display circuitry 37 displays the photon
counting CT image reconstructed in step S3.
[0093] The description of an example of the operation of photon
counting CT imaging according to this embodiment will be ended
here.
[0094] As described above, the photon counting CT apparatus
according to this embodiment includes the X-ray generation system
13, the X-ray detection system 15, the response function memory
circuitry 331, and the reconstruction circuitry 33. The X-ray
generation system 13 generates X-rays. The X-ray detection system
15 includes the X-ray detector 151 that detects the X-rays
generated by the X-ray generation system 13 and transmitted through
the subject S, and the data acquisition circuitry 153 that acquires
count data concerning the count number of X-rays detected based on
the output signal from the X-ray detector 151 for a plurality of
energy bands. The response function memory circuitry 331 stores
data of a response function that associates incident X-rays on the
X-ray detection system 15 with a detector response characteristic.
By executing the X-ray absorption amount calculation module 335,
the reconstruction circuitry 33 calculates the X-ray absorption
amount of each of a plurality of base substances based on count
data concerning a plurality of energy bands, the energy spectrum of
incident X-rays generated by the X-ray generation system 13, and
the response function readout from the response function memory
circuitry 331.
[0095] As described above, in this embodiment, the X-ray absorption
amount of each base substance is decided in considerate on of the
response function representing the detector response
characteristic. Hence, the decided X-ray absorption amount not
affected by the detector response characteristic, and its accuracy
improves as compared to an X-ray absorption amount calculated
without considering the detector response characteristic. The
response function according to this embodiment expresses not only a
main peak but also an escape peak or the energy of the
k-characteristic X-rays of a scintillator material. This makes it
possible to more accurately calculate the X-ray absorption amount
of a base substance. Hence, the accuracy of a photon counting CT
image reconstructed based on the X-ray absorption amount also
improves.
[0096] Thus, according to this embodiment, even a detector having a
low energy resolving power can accurately discriminate a
substance.
[0097] While certain embodiments have been described, these
embodiments have been presented by way of example only, and are not
intended to limit the scope of the inventions. Indeed, the novel
embodiments described herein may be embodied in a variety of other
forms; furthermore, various omissions, substitutions and changes in
the form of the embodiments described herein may be made without
departing from the spirit of the inventions. The accompanying
claims and their equivalents are intended to cover such forms or
modifications as would fall within the scope and spirit of the
inventions.
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