U.S. patent application number 17/144679 was filed with the patent office on 2021-07-29 for method and system for high-resolution x-ray detection for phase contrast x-ray imaging.
The applicant listed for this patent is KA IMAGING INC.. Invention is credited to Karim S. KARIM, Christopher C. SCOTT.
Application Number | 20210231588 17/144679 |
Document ID | / |
Family ID | 1000005510492 |
Filed Date | 2021-07-29 |
United States Patent
Application |
20210231588 |
Kind Code |
A1 |
KARIM; Karim S. ; et
al. |
July 29, 2021 |
METHOD AND SYSTEM FOR HIGH-RESOLUTION X-RAY DETECTION FOR PHASE
CONTRAST X-RAY IMAGING
Abstract
A phase contrast X-ray imaging system for imaging an object
including an X-ray source; and an X-ray detector having a 25 micron
or less pixel pitch; wherein a distance between the X-ray source
and the object is less than or equal to 10 cm. The X-ray detector
further includes at least one single direct conversion layer to
acquire at least one phase contrast edge-enhancement image.
Inventors: |
KARIM; Karim S.; (Waterloo,
CA) ; SCOTT; Christopher C.; (Waterloo, CA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
KA IMAGING INC. |
Waterloo |
|
CA |
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|
Family ID: |
1000005510492 |
Appl. No.: |
17/144679 |
Filed: |
January 8, 2021 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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16050354 |
Jul 31, 2018 |
10914689 |
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17144679 |
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62597622 |
Dec 12, 2017 |
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62573759 |
Oct 18, 2017 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61B 6/484 20130101;
G01N 23/041 20180201; A61B 6/482 20130101; A61B 6/5205 20130101;
A61B 6/542 20130101 |
International
Class: |
G01N 23/041 20060101
G01N023/041; A61B 6/00 20060101 A61B006/00 |
Claims
1. A phase contrast X-ray imaging system for imaging an object
comprising: an X-ray source; and a X-ray detector having a pixel
pitch less than 25 microns; wherein the X-ray detector includes at
least one single direct conversion layer to acquire at least one
phase contrast edge-enhancement image; and wherein a focal spot of
the X-ray source is less than or equal to 10 .mu.m.
2. The phase contrast X-ray imaging system of claim 1 wherein the
at least one single direct conversion layer comprises a
photoconductor layer.
3. The phase contrast X-ray imaging system of claim 2 wherein the
photoconductor layer comprises amorphous selenium, silicon, cadmium
zine telluride (CdZnTe), cadmium telluride (CdTe), mercury iodide
(HgI2), lead oxide (PbO) or scintillator infused organic
photoconductors.
4. The phase contrast X-ray imaging system of claim 1 wherein the
X-ray detector includes at least three direct conversion
layers.
5. The phase contrast X-ray imaging system of claim 1 wherein the
X-ray source comprises micro-focus X-ray tubes.
6. The phase contrast X-ray imaging system of claim 5 wherein the
micro-focus X-ray tubes comprise metal jet X-rays.
7. A method of phase contrast X-ray imaging comprising: placing an
X-ray source a distance R1 away from an object to be imaged;
placing an X-ray detector a distance R2 away from the object to be
imaged; directing a polychromatic beam at the object via the X-ray
source; detecting the X-ray photons via the X-ray detector; and
acquiring at least one phase contrast edge-enhancement image;
wherein the X-ray detector includes pixels having a pitch size less
than 25 microns.
8. The method of claim 7 wherein the directing a polychromatic beam
at the object via the X-ray source comprises: directing the
polychromatic beam from a metal jet X-ray source.
9. The method of claim 7 wherein the acquiring at least one phase
contrast edge-enhancement image comprises: detecting the at least
one phase contrast edge-enhancement via a direct conversion
layer.
10. The method of claim 9 wherein the direct conversion layer
comprises amorphous selenium, silicon, cadmium zine telluride
(CdZnTe), cadmium telluride (CdTe), mercury iodide (HgI2), lead
oxide (PbO) or scintillator infused organic photoconductors.
11. A phase contrast X-ray imaging system for imaging an object
comprising: an X-ray source; and a X-ray detector having a pixel
pitch less than 25 micron and at least one direct conversion layer
to acquire at least one phase contrast edge-enhancement image.
12. The phase contrast X-ray imaging system of claim 11 wherein the
at least one single direct conversion layer comprises a
photoconductor layer.
13. The phase contrast X-ray imaging system of claim 12 wherein the
photoconductor layer comprises amorphous selenium, silicon, cadmium
zine telluride (CdZnTe), cadmium telluride (CdTe), mercury iodide
(HgI2), lead oxide (PbO) or scintillator infused organic
photoconductors.
14. The phase contrast X-ray imaging system of claim 11 wherein the
X-ray detector includes at least three direct conversion
layers.
15. The phase contrast X-ray imaging system of claim 11 wherein the
X-ray source comprises micro-focus X-ray tubes.
16. The phase contrast X-ray imaging system of claim 15 wherein the
micro-focus X-ray tubes comprises metal jet X-rays.
17. The phase contrast X-ray imaging system of claim 11 wherein the
at least one contrast edge-enhancement image is used to form a
computed tomography image.
18. The phase contrast X-ray imaging system of claim 1 wherein the
at least one contrast edge-enhancement image is used to form a
computed tomography image.
19. The method of claim 7 further comprising: generating a computed
tomography image using the at least one contrast edge-enhancement
image.
20. The phase contrast X-ray imaging system of claim 17 wherein the
computer tomography image is a phase contrast edge-enhancement
computed tomography image.
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This application claims the benefit of priority of U.S.
Provisional Patent Applications No. 62/573,759 filed Oct. 18, 2017
and 62/597,622 filed Dec. 12, 2017 and is a continuation of U.S.
application Ser. No. 16/050,354 filed Jul. 13, 2018 which are all
hereby incorporated by reference.
FIELD OF THE DISCLOSURE
[0002] The disclosure is generally directed at X-ray imaging and,
more specifically, at a method and system for a high-resolution
X-ray detection for phase contrast imaging.
BACKGROUND OF THE DISCLOSURE
[0003] X-ray imaging has far-reaching applications in visualizing
objects using contrast provided by the heterogenous x-ray
absorption of their composition. Naturally, the utility of this
dominant paradigm of x-ray imaging diminishes if the penetrating
power of x-rays effectively make the object transparent. Such is
often the case for soft biological tissues or other low-density
materials such as plastics. In this context, we recall from optics
that electromagnetic waves have both an amplitude and a phase
associated with them. As x-rays penetrate the object, information
is not only encoded in the amplitude due to absorption, but also in
the phase due to refraction. This is analogous to a lens in optics,
where it is essentially transparent, however the refraction of
visible light encodes the shape of the lens. X-ray phase contrast
imaging (XPC) comprises methods of extracting phase information
from the x-ray intensity pattern detected by the detector.
[0004] The more practical solutions proposed to date for XPC
involve the use of multiple X-ray gratings and interferometry
techniques (i.e. Talbot Lau) which reduce the dose efficiency,
worsen spatial resolution, and increase cost and complexity of the
imaging chain making the entire system bulky and not suitable for
low-cost compact applications (e.g. benchtop XPC). All but the
simplest method, propagation-based XPC (PB-XPC), requires
additional apparatus.
[0005] Using PB-XPC, the ability to retrieve phase information,
that is to detect the very small refraction angles of x-rays, falls
entirely on the capabilities of the x-ray source. To date, PB-XPC
is a common technique used at synchrotron facilities where the
following three critical requirements are simultaneously met for
PB-XPC: (1) monochromatic X-rays to facilitate ease of image
reconstruction, (2) spatially coherent X-rays that can provide a
correlated wave-field from which to detect phase changes and (3)
since spatial coherence is proportional to the source-to-object
distance, a high flux of X-rays is necessary because the object is
placed far from the source and X-ray intensity is inversely
proportional to the square of the distance. Although the PB-XPC
technique has proven to be useful, it is practically limited to use
at synchrotron facilities. Thus, there is still a need for a
compact and fast X-ray phase contrast imaging system for home lab
life sciences, health and scientific imaging, and non-destructive
test applications that is based on PB-XPC but does not require a
synchrotron source to successfully image low density materials at
low X-ray exposures.
[0006] Therefore, there is provided a novel method and system for
high-resolution X-ray detection for phase contrast imaging
SUMMARY OF THE DISCLOSURE
[0007] In one aspect of the disclosure, there is provided a phase
contrast X-ray imaging system for imaging an object including an
X-ray source; and an X-ray detector having a 25 micron or less
pixel pitch; wherein a distance between the X-ray source and the
object (R.sub.1-1) is less than or equal to 10 cm.
[0008] In another aspect, R.sub.1-1 is a distance between a source
focal point of the X-ray source and an object plane of the object.
In a further aspect, a distance between the X-ray detector and the
object (R.sub.2-1) is greater than 0 cm. In yet another aspect,
R.sub.2-1 is a distance between an object plane of the object and a
detector plane of the X-ray detector. In an aspect, R.sub.2-1 is
less than or equal to 200 cm.
[0009] In a further aspect, the system further includes a second
X-ray source; and a second X-ray detector; wherein a distance
between the second X-ray source and the object (R.sub.1-2) is less
than or equal to 10 cm. In another aspect, a distance between the
second X-ray detector and the object (R.sub.2-2) is greater than 0
cm. In another aspect, the X-ray source and the second X-ray source
shine X-ray beams towards the object in non-parallel directions. In
yet a further aspect, the X-ray source and the second X-ray source
shine X-ray beams towards the object in perpendicular directions.
In an aspect, a focal spot of the X-ray source is <30 .mu.m. In
another aspect, the X-ray detector is a multi-layer X-ray detector.
In yet another aspect, the multi-layer X-ray detector includes
direct conversion layers. In another aspect, the multi-layer X-ray
detector includes direct and indirect conversion layers. In yet
another aspect, the multi-layer X-ray detector includes indirect
conversion layers.
[0010] In another aspect of the disclosure, there is provided a
method of phase contrast X-ray imaging including placing an X-ray
source a distance R.sub.1 away from an object to be imaged; placing
an X-ray detector a distance R.sub.2 away from the object to be
imaged; directing a polychromatic beam at the object via the X-ray
source; and detecting the X-ray photons via the X-ray detector;
wherein the X-ray detector includes pixels having a size less than
or equal to 25 microns; and wherein R.sub.1 is less an 10 cm. In
another aspect, R.sub.2 is between 0 cm and 200 cm.
[0011] In another aspect of the disclosure, there is provided a
phase contrast X-ray imaging system for imaging an object including
an X-ray source; and an X-ray detector; wherein a distance between
the X-ray source and the object (R.sub.1) is less than or equal to
10 cm; and wherein a distance between the X-ray detector and the
object (R.sub.2) is between 0 and 200 cm.
[0012] In another aspect, R.sub.1 is measured between an output of
the X-ray source and an object plane of the object. In yet another
aspect, R.sub.2 is measured between a detector plane of the X-ray
detector and an object plane of the object.
BRIEF DESCRIPTION OF THE DRAWINGS
[0013] Embodiments of the present disclosure will not be described,
by way of example only, with reference to the attached Figures.
[0014] FIG. 1 is a schematic diagram of a propagation-based X-ray
phase contrast imaging system;
[0015] FIG. 2 is a schematic diagram of a cross-section of the
direct-conversion x-ray detector;
[0016] FIG. 3 is a photograph of a digital X-ray detector for use
in the system of FIG. 1;
[0017] FIG. 4a is a graph showing DQE vs spatial frequency using
the X-ray detector of FIG. 3;
[0018] FIG. 4b is a graph showing DQE vs spatial frequency using
known X-ray detectors;
[0019] FIG. 5a is an X-ray image of a bell-pepper seed absorption
image with phase contrast reduced;
[0020] FIG. 5b is an X-ray image of a bell-pepper seed absorption
image with phase contrast;
[0021] FIG. 6 is a schematic diagram of a multilayer detector
having layers 1 to N;
[0022] FIG. 7 is a schematic diagram of a first embodiment of a
system configuration to obtain multi-energy X-ray images and phase
contrast images simultaneously;
[0023] FIG. 8 is a graph showing penetration depth of x-ray photons
in amorphous selenium photoconductor material; and
[0024] FIG. 9 is a flowchart outlining a method of phase contrast
X-ray imaging.
DETAILED DESCRIPTION OF THE DISCLOSURE
[0025] The disclosure is directed at a method and system for a
high-resolution X-ray detection for phase contrast imaging. In one
embodiment, the system includes an X-ray source and an X-ray
detector with a pixel pitch of less than or equal to 25 microns.
The X-ray source is preferably located a distance R.sub.1 from an
object plane while the X-ray detector is preferably located a
distance R.sub.2 from the object plane.
[0026] Turning to FIG. 1, a schematic diagram of a system for
high-resolution X-ray detection for phase contrast imaging is
shown. The system may be seen as a propagation-based X-ray phase
contract imaging system. In one embodiment, the system enables
propagation-based X-ray phase contrast imaging (PB-XPC) in a
compact, fast manner by approaching PB-XPC from a source and
detector perspective. The system 10 includes an X-ray source 12
that directs X-rays (such as in the form of a polychromatic beam
14) towards an object 16 that is being imaged. The system further
includes a detector 18, located on a side opposite the X-ray source
with respect to the object 16) to receive, or detect, the X-rays
that pass through the object 16 through free-space propagation. In
a preferred embodiment, the X-ray source 12 is a standard
laboratory micro-focus source and the X-ray detector 18 is a very
high resolution and dose efficient X-ray detector having a pixel
pitch of less than or equal to 25 microns.
[0027] As shown in FIG. 1, an output plane 20 of the focal spot of
the X-ray source 12 is located a distance R.sub.1 from the object
plane 22 while an image plane 24 of the X-ray detector 18 is a
distance R.sub.2 from the object plane 22. By selecting a
corresponding pixel pitch (preferably less than or equal to 25
microns), an optimal (or preferred) R.sub.1 (which can be seen as
an X-ray source focal spot to object plane/source to object
distance) and an optimal (or preferred) R.sub.2 (which may be seen
as an object plane to detector image plane/object to detector
distance) may be selected to achieve, fast, dose efficient PB-XPC
using a benchtop device. In one embodiment, the selection of the
pixel pitch may be based on the X-ray refraction angle of the X-ray
leaving the object (calculated from the complex refractive index)
and the propagation distance R.sub.2. In a preferred embodiment, a
small R.sub.2 is more desirable, leading to a deviation of the
X-ray on the that is detectable by a detector having pixels with a
small pixel pitch (such as less than or equal to 25 microns).
[0028] As was experienced during experiments, the system may detect
the minute (in the range of 10.sup.-5-10.sup.-4 rad) X-ray
refraction associated with phase changes encoded by the object
16.
[0029] In one preferred embodiment, the X-ray source 12 may be a
standard low-power (8 W) laboratory micro-focus source with a focal
spot size of 5 to 9 .mu.m. The focal spot size is the size of the
X-ray source electron beam that contacts the anode target materials
e.g. tungsten or molybdenum, which then produces X-rays that
propagate to the object 16 and subsequently to the detector 18). In
current medical imaging solutions, the focal spot size is 0.3 to 1
mm. When the focal spot is small (such as between 5 to 9 .mu.m),
the penumbral blur from the extent of the focal spot is minimized
or reduced such that that the X-ray source 12 does not limit
spatial resolution within the system 10. Given the aim to detect
phase changes due to the object 16, a coherent or partially
coherent incident beam is necessary or preferred. The lateral
coherence length is proportional to the source-to-object distance,
R.sub.1, and inversely proportional to the focal spot size. That
is, a smaller focal spot results in a partially coherent beam with
a smaller R.sub.1 distance, or in other words, a more compact
system.
[0030] One challenge is that a small focal spot in a traditional
fixed anode (i.e. not a costly liquid-metal jet source), the
micro-focus source results in low power output due to the heat load
on the object. This limitation is a key challenge in obtaining a
phase contrast image in both a short time and at low x-ray
exposures (e.g. to minimize or reduce radiation damage to objects
such as, but not limited to, biological samples).
[0031] Turning to FIG. 2, a schematic cross-section of an X-ray
detector is shown. In the current disclosure, the detector is
preferably a high-resolution x-ray detector based using a direct
conversion photoconductor and complementary metal-oxide
semiconductor (CMOS) pixel electronics having a pixel pitch of less
than or equal to 25 microns.
[0032] As shown in FIG. 2, the X-ray detector 18 includes a bottom
CMOS layer 30 with a plurality of small sized pixels 32. In the
current disclosure, the pixel pitch of each of the pixels 32 is
less than or equal to twenty-five (25) microns. The detector 18
further includes a stability/blocking layer 34, a photoconductor
layer 36, a blocking layer 38 and an electrode layer 40. The
detector 18 may further include a set of bond pads 42 that are used
to enable an electrical connection for control/data signals.
[0033] In one embodiment, the photoconductor layer 36 is an
amorphous selenium (a-Se) photoconductor layer 36. In this
embodiment, the blocking layers 34 and 38 on either side of the
a-Se photoconductor layer 36 may be used to improve mechanical
stability of the detector 18 and/or to reduce the dark current
during operation of the detector 18 at high electric fields. In
another embodiment, the detector 18 may include only one or none of
the blocking layers 34 or 38.
[0034] In another embodiment, the stability/blocking layer 34 may
be a polyimide layer that may function as both, an
anticrystallization layer and as a blocking contact on the bottom
of the photoconductor layer 36. In another embodiment, the blocking
layer 38 may be a parylene layer that functions as a blocking
contact for the photoconductor layer 36. A contact layer between
the photoconductor layer 36 and the stability/blocking layer may
also be, but is not limited to, a p-type layer (such as As-doped
selenium) or other soft polymer materials. A contact layer between
the photoconductor layer 36 and the blocking layer 38 may also be,
but is not limited to, a n-type layer such as alkali-metal-doped
selenium or cold deposited selenium, or other known organic and
inorganic hole blocking layers. Although the previous discussion
relates to a direct conversion X-ray detector, other
high-resolution detector technologies, such as indirect conversion
detectors, or a combination of direct conversion and indirect
conversion X-ray detectors are contemplated.
[0035] In direct conversion X-ray detectors, amorphous selenium,
silicon, CdZnTe, CdTe, HgI.sub.2, PbO, and scintillator infused
organic photoconductors such as perovskite integrated with CMOS or
thin-film-transistor (TFT) pixel arrays may be used for the
photoconductor layer 36. With indirect conversion X-ray detectors,
CsI, LaBr.sub.3, and pixelated GOS or CsI scintillators integrated
CMOS or TFT pixel arrays are may be used.
[0036] Excluding x-ray obliquity, which affects both indirect and
direct conversion detectors, the thickness of the direct conversion
photoconductor within the X-ray detector does not have the same
trade-off with spatial resolution as an indirect conversion
photoconductor because a large applied electric field transports
the X-ray generated charge carriers with negligible lateral
diffusion.
[0037] One advantage of the disclosure is the use of a very fine,
or small, pixel pitch, high dose efficiency direct conversion X-ray
detector to work in conjunction with the micro-focus source 12 for
the PB-XPC approach.
[0038] Current X-ray indirect-detection technology exhibits a
tradeoff between spatial resolution and dose efficiency. The
scintillator material used to convert x-rays to optical photons for
detection by a pixelated matrix of photodiodes results in increased
optical scatter with thickness. Thicker scintillators absorb more
photons but also lead to increased light scattering while thin
scintillators preserve resolution by limiting scatter but absorb
fewer photons and are dose inefficient reducing the detective
quantum efficiency (DQE). Moreover, trying to visualize very fine
features with lower spatial resolution detectors requires a large
magnification factor which, when coupled with micro focal spot (and
thus, lower power) X-ray sources additionally leads to longer scan
times and dose.
[0039] Turning to FIG. 3, a photograph of one embodiment of a pixel
pitch imager is shown. The pixel pitch imager of FIG. 3 is a 5.5
um.times.6.25 um pixel pitch imager. Through experimentation, the
dose efficiency measurements were around 10.times. better than
current systems and projected results that may be up to 100.times.
better than current detectors by using pixels having a size less
than or equal to 25 microns. Imaging time can be further reduced by
using high output micro-focus X-ray tubes (e.g. metal jet X-ray) as
the X-ray source, however, use of a high dose efficiency detector
helps further reduce imaging time (e.g. for high throughput
industrial applications) and more importantly, to minimize or
reduce further radiation damage to sensitive biological tissue,
especially in life sciences and medical applications.
[0040] Furthermore, in the micrograph of FIG. 3, the pixel imager,
or hybrid a-Se/CMOS digital X-ray detector, the overall chip
dimensions are 1.8.times.3.0 mm.sup.2. The a-Se/CMOS hybrid
structure is visible with a biasing probe for application of
positive high voltage to the gold electrode.
[0041] In FIG. 4a, which reflect results/measurements using the
X-ray detector of the disclosure, the DQE calculated for the 70 kVp
spectrum using the measured modulation transfer function (MTF) and
measured noise power spectrum (NPS) are shown. The results in the
20-60 cycles/mm range exceed all other previously reported X-ray
detector DQE results. FIG. 4b shows a modeled DQE at 70 kVp for an
absorption-optimized a-Se photoconductor layer with a thickness of
1000-.mu.m assuming no focal spot blur and 100 e.sup.- RMS read-out
noise. With optimized X-ray absorption, the DQE is very high (above
0.5 or 50%) in the 20-60 cycles/mm range. For the graph of FIG. 4b,
the photoconductor thickness for the modelled detector is 1000
microns while the photoconductor thickness for the detector of FIG.
4a was 56 microns.
[0042] Using the phase contrast X-ray system of the disclosure, the
added detail due to phase contrast is demonstrated in FIGS. 5a and
5b. The hook was used to suspend the bell pepper seed which served
as the object being imaged. In the case of this phase contrast
image, the source-to-detector distance was 26 cm (sum of
R.sub.1+R.sub.2), allowing the images to be taken in a few seconds
compared to the minutes and hours commonly reported for current
phase contrast systems. As such, the system of the disclosure may
be seen as a highly compact, fast, low dose PB-XPC systems. In this
experiment, R.sub.1 was less than 10 cm for the images captured
(with R.sub.2 greater than 0 cm). The R.sub.1 values used in the
system of the disclosure are in direct contrast to current PB-XPC
systems which teach away from using R.sub.1 values of <10
cm.
[0043] Using the system of the disclosure, phase contrast images
were achieved with R.sub.1 values of <10 cm for a range of
R.sub.2 values (e.g. between 0 and 200 cm) and pixel sizes of less
than or equal to 25 microns. In one embodiment, pixels sizes less
than 10 microns are contemplated.
[0044] In simulations, a source focal spot of <30 .mu.m was
shown to be suitable for phase contrast imaging although a focal
spot of <10 .mu.m is preferable for sharper images and a more
compact system.
[0045] Turning to FIG. 6, a diagram of another embodiment of an
X-ray detector for use with the system of the disclosure is shown.
The X-ray detector 18 of FIG. 6 may be seen as a multi-layer
detector and may enable a compact X-ray imaging system that
acquires both: multi-spectral (e.g. dual energy spectral X-ray
data) as well as a phase contrast image (including phase retrieval)
simultaneously.
[0046] In the current embodiment, the X-ray detector 18 includes a
set of conversion layers 100 (seen as Conversion layer 1,
Conversion layer 2, . . . Conversion layer N (where N is any
number)) a set of substrate layers 102 and a set of X-ray filters
104. Different design/structure of the conversion layers 100,
substrate layers 102 and X-ray filters 104 are contemplated and
FIG. 6 provides one such example structure. As will be appreciated,
the simplest implementation of such a multi-layer detector would
include two stacked conversion layers 100 with an intermediate
mid-filter 104. An improved approach could use three stacked
conversion layers with the middle conversion layer acting as a
mid-filter. As will be understood, each of the conversion layers is
associated with a set of pixels having a size of less than or equal
to 25 microns. With N conversion layers and N set of pixels, N
unique data sets may be simultaneously obtained or generated at a
low object dose i.e. multi-spectral, phase contrast, along with an
original attenuation image.
[0047] In the Fresnel region, the "transport of intensity equation"
(TIE) implies that contrast from intensity variations at the image
plane is proportional to the propagation distance from the object
plane and the spatial gradient of the phase distribution in the
object plane. This differential phase contrast results in an
"edge-enhancement" effect due to phase changes being most abrupt at
the edges of the object where there is a rapid change in the
refractive index. Although the use of PB-XPC X-ray imaging results
in increased contrast at object boundaries for better detectability
of materials with poor x-ray absorption, the relationship between
the physical geometry of the object and its visualization in the
image plane is more complicated.
[0048] Specifically, the boundaries in the image may not correspond
exactly to boundaries in the object. To restore quantitative
boundary information in the image, a "phase retrieval"
reconstruction is typically required to be performed. One method
for phase retrieval is a "direct approach" by solving the
deterministic TIE for x-ray intensity and phase information in the
object plane. Being non-iterative and numerically efficient this
method is viable for use in projection imaging and for 3D
micro-CT.
[0049] The TIE, for a single wavelength, includes one known
variable (intensity in the image plane) and two unknown variables
(intensity and phase in the object plane). In the case of a pure
phase (i.e. no absorption) or homogenous object and monochromatic
radiation, the solution to the TIE is relatively straightforward.
For this case, in the geometric optics approximation, the intensity
and phase in the object plane are related and a unique solution to
the TIE can be obtained from a single measurement in the image
plane or alternately, a single image acquisition.
[0050] For general inhomogeneous objects (i.e. the more practical
situation) with uncorrelated absorption and refraction properties,
at least two measurements at different image planes or different
radiation wavelengths are required to solve the system of
equations. This requirement poses a challenge for radiation dose
sensitive (life sciences or medical) or even high throughput (e.g.
real-time) applications where the time taken to move the detector
to acquire the two measurements (i.e. images) necessary for phase
retrieval is prohibitive. As such, the system of the disclosure
allows for multiple images to be retrieved with a lower dose
exposure for the object. Moreover, most practical applications
(e.g. biomedical clinical imaging or even in industrial inspection)
require the use of commonly available polychromatic x-ray sources,
which makes obtaining the conventional TIE solution problematic
since it inherently assumes a monochromatic source.
[0051] To overcome the above challenges of obtaining at least two
measurements to solve the TIE with monochromatic and/or
polychromatic sources, the multilayer (i.e. stacked) X-ray detector
of FIG. 6 may be used to simultaneously capture multiple images at
different image planes with adaptable X-ray spectra for PB-XPC. A
multilayer detector typically includes a plurality of stacked x-ray
conversion layers on optional substrates with optional intermediate
x-ray filter materials (such as schematically shown in FIG. 6),
where critically, each conversion layer captures information in a
different image plane.
[0052] Each conversion layer can be a direct conversion layer (such
as the proposed fine pitch a-Se direct conversion X-ray detector)
or an indirect conversion layer. In a direct conversion layer, an
X-ray semiconductor (e.g. amorphous selenium, silicon, PbO, HgI2,
CdZnTe, CdTe, organic semiconductor with nanoparticles, etc.)
converts incident X-ray photons directly into electronic charge.
The X-ray semiconductor can be optionally paired with a readout
electronics plane (e.g. thin film transistor array, CMOS pixel
array) that contains an active matrix array of readout pixels
(transistors and/or storage capacitor). In certain cases, the X-ray
semiconductor and readout electronics plane are both part of the
X-ray conversion layer.
[0053] In an indirect X-ray conversion layer, the scintillator
material (e.g. GOS, CsI, NaI, CaWO4, LYSO, etc.) is used to convert
incident X-ray photons into optical photons, which are then
detected by an underlying pixelated photosensitive readout
electronics plane. The photosensitive readout electronics plane
could be a large area active matrix array of pixels (e.g.
containing a photodiode with thin film transistors or a photodiode
with an active pixel sensor) made of a variety of materials
including large area thin film inorganic (e.g. amorphous silicon,
metal oxide, LTPS, continuous grain silicon, crystalline silicon)
or even organic semiconductors. In this embodiment, the
scintillator and photosensitive readout electronics can both be
part of the X-ray conversion layer.
[0054] Due to the greater penetration depth of higher energy
photons relative to lower energy photons (e.g. see FIG. 8 for
penetration depth in amorphous selenium semiconductor), a single
x-ray exposure results in each X-ray conversion layer acquiring an
image with a different x-ray spectrum. The X-ray spectra can be
controlled using the thickness of each conversion layer (i.e. the
semiconductor layer in direct conversion or the scintillator layer
in indirect conversion) and/or the filter layer. Characterization
of the spectra (without an object) may be necessary for phase
retrieval.
[0055] In one embodiment, the penetration depth is equal to the
reciprocal of the X-ray attenuation coefficient and corresponds to
the depth within a material that the x-ray intensity reduces to
.about.37% of its initial value. The discontinuity at .about.12.7
keV is due to photoelectric absorption.
[0056] Filter materials can range from common metal mid-filters,
such as aluminum and copper. If an additional X-ray conversion
layer is used as the filter, then, in this case, there would be
three X-ray conversion layers stacked on top of each other. In
principle, at least two X-ray conversion layers are necessary but
additional layers can be stacked as necessary to obtain additional
spectral separation, which could improve phase retrieval by
allowing the use of more accurate reconstruction formulae.
[0057] Even further spectral separation could be obtained by
modulating the X-ray semiconductor thickness in any given direct
X-ray conversion layer on a pixel by pixel basis or alternately,
modulating the scintillator thickness in any given indirect X-ray
conversion layer on a pixel by pixel basis. By modulating the
thickness of the X-ray conversion layer at the pixel level, spatial
resolution can be a trade-off to obtain extra spectral separation
even in a single layer.
[0058] Using very small pixel pitch dimensions (as with our fine
pixel pitch detector having pixel sizes less than or equal to 25
microns) in each conversion layer can further improve performance
by detecting the small refraction angle of x-rays (which is
necessary for phase contrast) at shorter propagation distances from
object plane to image plane. X-ray intensity (and therefore
signal-to-noise ratio) decreases with the inverse square of
propagation distance, so reducing propagation distance can lower
dose as well as potentially speed up phase retrieval compared to
other propagation-based methods or other phase contrast imaging
modalities (e.g. grating based.)
[0059] In another embodiment, to obtain both multi-spectral and
phase retrieval data for PB-XPC, the system may include two
different X-ray sources in conjunction with two fine-pitch single
layer X-ray detectors that are operating in different planes as
schematically shown in FIG. 7. As will be understood, a fine-pitch
single layer X-ray detector is one with pixels having a size less
than or equal to 25 microns.
[0060] As shown in FIG. 7, the system includes a first X-ray source
150 that directs a polychromatic beam towards an object 152 that is
then detected by a first X-ray detector 154. The system further
includes a second X-ray source 156 that directs a polychromatic
beam towards the object 152 that is then detected by a second X-ray
detector 158. In one embodiment, the distance between the first
X-ray source 150 and the object plane (R1.sub.D1 or R.sub.1-1) and
the distance between the second X-ray source 156 and the object
plane (R1.sub.D2 or R.sub.1-2) may be set to the same value while
the distance between the object plane and the image plane of the
first X-ray detector 154 (R2.sub.D1 or R.sub.1-2) and the distance
between the image plane of the second X-ray detector 158 and the
object plane (R1.sub.D2 or R.sub.2-2) may be set to different
values. The two set of X-ray source and X-ray detector pairs allow
the system to obtain multiple two-dimensional (2D) images from the
first and second X-ray detectors. In an alternate embodiment, the
beams of the first X-ray source and the second X-ray source shine
X-ray are directed towards the object in non-parallel directions.
In another embodiment, the beams of the first X-ray source and the
second X-ray source are directed towards the object in
perpendicular directions.
[0061] In both embodiments where multiple images are generated or
detected, they may then be combined in any known methodologies to
obtain a single overall image (if required) using reconstruction
algorithms.
[0062] One advantage of the system of FIG. 7 is that the X-ray
spectrum from the first X-ray source 150 and the X-ray spectrum
from the second X-ray source 156 may be defined independently of
the first X-ray detector 154 and the second X-ray detector 158
leading to additional simplicity in the reconstruction algorithms.
As before, the system configuration of FIG. 7 may enable
acquisition of phase contrast images, phase retrieval,
multi-spectral images and conventional attenuation images in a
single scan. To obtain a three-dimensional (3D) image, either the
object or the source/detector pairs can be rotated to obtain
multiple projections for reconstruction or further X-ray
source/X-ray detector pairs may be used.
[0063] Turning to FIG. 9, a flowchart outlining a method of phase
contrast imaging is shown. Initially, an X-ray source is placed a
distance R1 away from the object being imaged (900). This distance
is preferably less than 10 cm and, in one embodiment, is measured
from the focal spot of the X-ray source to the object plane of the
object. An X-ray detector is then placed a distance R.sub.2 from
the object (902) on a side of the object opposite the location of
the X-ray source. This distance is preferably between 0 cm and 200
cm and, in one embodiment, is measured from the object plane to a
detector plane.
[0064] The X-ray source then directs a polychromatic beam towards
the object (904). The resulting photons are then detected by the
X-ray detector via its set of pixels that are sized to be less than
or equal to 25 microns (906). If necessary, further X-ray source
and X-ray detector pairs may be placed (908) around the object to
obtain multiple images with a lower radiation dose.
[0065] While the current disclosure has been directed at a compact
phase contrast X-ray detector with direct conversion selenium-CMOS
detectors, other direct conversion materials such as HgI.sub.2,
CZT, TIBr, and silicon can be employed in place of selenium and the
CMOS pixels could be replaced by poly-Si, metal-oxide, or common
II-VI or III-V semiconductors. Moreover, high-resolution
indirect-conversion X-ray detectors (e.g. with thin scintillators,
or pixelated scintillators) can also be employed albeit likely with
lower dose efficiency than direct conversion detectors.
Micro-computed-tomography (microCT) is also possible with this
system by adding a rotational stage (or creating a rotating gantry)
for generating multiple x-ray projection images of the object from
different perspectives, and CT reconstruction software.
[0066] In addition to providing fast imaging in a compact system,
the system of the disclosure also has a significant benefit for
micro-anatomical imaging to visualize greater level of detail and
avoid damaging DNA by using less X-ray radiation to acquire an
image. As an example, since detailed knowledge of genes and the
ability to control gene expression is available in mice and rats,
the ability to quantitate the impact of highly targeted genetic
manipulations on organ structure and function using phase contrast
micro-CT could help answer how genes link to whole body
pathophysiology. The combination of better visualization of soft
tissue using phase contrast X-ray and high detector dose efficiency
can fundamentally advance genomics by allowing high resolution,
non-invasive and non-destructive imaging in live, intact animals
and plants, tissues, and even single cells--tasks that are not
possible using other techniques. Similar advantages exist for other
scientific and non-destructive imaging applications for example,
imaging agricultural products, plastics, polymers and various
nano-composite materials and glasses.
[0067] In the preceding description, for purposes of explanation,
numerous details are set forth in order to provide a thorough
understanding of the embodiments. However, it will be apparent to
one skilled in the art that these specific details may not be
required. In other instances, well-known structures may be shown in
block diagram form in order not to obscure the understanding. For
example, specific details are not provided as to whether elements
of the embodiments described herein are implemented as a software
routine, hardware circuit, firmware, or a combination thereof.
[0068] Embodiments of the disclosure or components thereof can be
provided as or represented as a computer program product stored in
a machine-readable medium (also referred to as a computer-readable
medium, a processor-readable medium, or a computer usable medium
having a computer-readable program code embodied therein). The
machine-readable medium can be any suitable tangible,
non-transitory medium, including magnetic, optical, or electrical
storage medium including a diskette, compact disk read only memory
(CD-ROM), memory device (volatile or non-volatile), or similar
storage mechanism. The machine-readable medium can contain various
sets of instructions, code sequences, configuration information, or
other data, which, when executed, cause a processor or controller
to perform steps in a method according to an embodiment of the
disclosure. Those of ordinary skill in the art will appreciate that
other instructions and operations necessary to implement the
described implementations can also be stored on the
machine-readable medium. The instructions stored on the
machine-readable medium can be executed by a processor, controller,
or other suitable processing device, and can interface with
circuitry to perform the described tasks.
[0069] The above-described embodiments are intended to be examples
only. Alterations, modifications and variations can be effected to
the particular embodiments by those of skill in the art without
departing from the scope, which is defined solely by the claims
appended hereto.
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