U.S. patent application number 10/212300 was filed with the patent office on 2002-12-26 for system for quantitative radiographic imaging.
This patent application is currently assigned to University of Massachusetts Medical Center. Invention is credited to Karellas, Andrew.
Application Number | 20020196899 10/212300 |
Document ID | / |
Family ID | 27031801 |
Filed Date | 2002-12-26 |
United States Patent
Application |
20020196899 |
Kind Code |
A1 |
Karellas, Andrew |
December 26, 2002 |
System for quantitative radiographic imaging
Abstract
A system for spectroscopic imaging of bodily tissue in which a
scintillation screen and a charged coupled device (CCD) are used to
accurately image selected tissue. An x-ray source generates x-rays
which pass through a region of a subject's body, forming an x-ray
image which reaches the scintillation screen. The scintillation
screen reradiates a apatial intensity pattern corresponding to the
image, the pattern being detected by a CCD sensor. The image is
digitized by the sensor and processed by a controller before being
stored as an electronic image. Each image is directed onto an
associated respective CCD or amorphous silicon detector to generate
individual electronic representations of the separate images.
Inventors: |
Karellas, Andrew; (Auburn,
MA) |
Correspondence
Address: |
THOMAS O. HOOVER, ESQ.
BOWDITCH & DEWEY, LLP
161 Worcester Road
P.O. Box 9320
Framingham
MA
01701-9320
US
|
Assignee: |
University of Massachusetts Medical
Center
Worcester
MA
|
Family ID: |
27031801 |
Appl. No.: |
10/212300 |
Filed: |
August 5, 2002 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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10212300 |
Aug 5, 2002 |
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09389760 |
Sep 2, 1999 |
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6445767 |
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09389760 |
Sep 2, 1999 |
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08469895 |
Jun 6, 1995 |
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6031892 |
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08469895 |
Jun 6, 1995 |
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08438800 |
May 11, 1995 |
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08438800 |
May 11, 1995 |
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07853775 |
Jun 2, 1992 |
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5465284 |
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07853775 |
Jun 2, 1992 |
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PCT/US90/07178 |
Dec 5, 1990 |
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PCT/US90/07178 |
Dec 5, 1990 |
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07446472 |
Dec 5, 1989 |
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5150394 |
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Current U.S.
Class: |
378/98.8 ;
348/E5.027; 348/E5.086 |
Current CPC
Class: |
A61B 6/505 20130101;
H04N 5/2253 20130101; A61B 6/145 20130101; A61B 6/4241 20130101;
A61B 6/4423 20130101; A61B 6/4233 20130101; G21K 1/10 20130101;
A61B 6/508 20130101; A61B 6/4035 20130101; A61B 6/405 20130101;
A61B 6/4291 20130101; A61B 6/06 20130101; A61B 6/4488 20130101;
G21K 4/00 20130101; A61B 6/4258 20130101; A61B 6/502 20130101; A61B
5/0091 20130101; H04N 5/32 20130101 |
Class at
Publication: |
378/98.8 |
International
Class: |
H05G 001/64 |
Claims
What is claimed:
1. An x-ray imaging device comprising: a support surface that
supports a region of interest of a patient to be imaged between an
x-ray source and an x-ray conversion device, the conversion device
having a first two-dimensional area and emitting a signal
correlated with the received radiation, said signal extending
across the first two-dimensional area; and a solid state device
having a two dimensional array of pixels and generating an
electronic representation of the region of interest, the signal
received at the array of pixels extending across a second
two-dimensional area.
2. The device of claim 1 wherein the first and second
two-dimensional areas have the same dimensions.
3. The device of claim 1 further comprising a display that displays
the electronic representation of the region of interest.
4. The device of claim 1 further comprising a frame holding the
source and the solid state device in fixed relation to each
other.
5. The device of claim 1 further comprising a data storage device
that electronically stores the electronic representation of the
region of interest.
6. The device of claim 1 further comprising a radiation source
positioned above the patient whose tissue is being examined.
7. The device of claim 1 wherein the solid state device comprises
an integrated silicon circuit.
8. The device of claim 1 wherein the array of pixels is larger than
512.times.512 pixels.
9. An x-ray imaging device comprising: a support surface that
supports a region of interest of a patient to be imaged between an
x-ray source and an x-ray detector, the detector including a solid
state device having a two dimensional array of pixels and
generating an electronic representation of the region of interest,
the signal received at the array of pixels; and a controller
connected to the detector to control readout of image data to an
electronic memory.
10. The device of claim 9 wherein the detector comprises a
cassette.
11. The device of claim 9 further comprising a display that
displays the electronic representation of the region of
interest.
12. The device of claim 9 further comprising a frame holding the
source and the solid state device in fixed relation to each
other.
13. The device of claim 9 further comprising a data storage device
that electronically stores the electronic representation of the
region of interest.
14. The device of claim 9 further comprising a radiation source
positioned above the patient.
15. The device of claim 9 wherein the solid state device comprises
an integrated silicon circuit.
16. The device of claim 9 wherein the array of pixels is larger
than 512.times.512 pixels.
17. The device of claim 9 wherein the detector includes an
amorphous silicon sensor.
18. The device of claim 9 further comprising a fiber optic coupler.
Description
CROSS REFERENCES TO RELATED APPLICATIONS
[0001] This application is a continuation of U.S. patent
application Ser. No. 09/389,760 filed Sep. 2, 1999 which is a
continuation of U.S. patent application Ser. No. 08/469,895 filed
Jun. 6, 1995, now U.S. Pat. No. 6,031,892 which is a
continuation-in-part of U.S. patent application Ser. No. 08/438,800
filed May 11, 1995 which is a continuation-in-part of U.S. patent
application Ser. No. 07/853,775 filed Jun. 2, 1992, now U.S. Pat.
No. 5,465,284, which is the U.S. National Phase of International
Application No. PCT/US90/07178, filed Dec. 5, 1990 and which is a
continuation-in-part of U.S. patent application Ser. No.
07/446,472, filed Dec. 5, 1989, now U.S. Pat. No. 5,150,394. The
entire contents of the above applications are incorporated herein
by reference in entirety.
BACKGROUND OF THE INVENTION
[0002] In recent years the use of radiological examining equipment
to make measurements of bone density in patients has continually
increased. In particular, the use of such equipment in diagnosing
and analyzing osteoporosis has become prevalent in the medical
community. Osteoporosis is characterized by the gradual loss of
bone mineral content or atrophy of skeletal tissue, resulting in a
corresponding overall decrease in average bone density. Such a
condition is common in elderly women and greatly increases the risk
of fracture or similar bone related injury.
[0003] The presently available techniques for the radiological
measurement of bone density utilize a rectilinear scanning
approach. In such an approach, a radiation source, such as a
radionuclide source or an x-ray tube, and a point detector are
scanned over a patient in a raster fashion. This scan results in an
image which has been derived from the point-by-point transmission
of the radiation beam through the bone and soft tissue of a
patient. The calculation of the bone-mineral concentration (the
"bone density") is usually performed by a dual energy approach.
[0004] The current rectilinear scanning approach is generally
limited by its long scanning time and its lack of good spatial
resolution. The poor spatial resolution results in an inability to
provide an image displaying high anatomical detail and which will
permit accurate determination of the area in the scan occupied by
bone. Moreover, the output of the x-ray source and the response of
the detector must be closely monitored in order to assure high
accuracy and precision.
SUMMARY OF THE INVENTION
[0005] In accordance with the present invention, a stationary bone
densitometry apparatus is provided for examining a subject's body.
A dual energy x-ray source directs a beam of x-ray radiation toward
the subject's body. The radiation is applied to the entire region
of the body being examined. A scintillation screen receives the
x-ray radiation passing through the body of the subject, and emits
radiation in the visible spectrum with a spatial intensity pattern
proportional to the spatial intensity pattern of the received x-ray
radiation.
[0006] A charge coupled device (CCD) then receives radiation from
the scintillation screen. This CCD sensor generates a discrete
electronic representation of the spatial intensity pattern of the
radiation emitted from the scintillation screen. A focusing element
between the screen and the CCD sensor focuses the scintillation
screen radiation onto the CCD sensor. To prevent ambient radiation
from reaching the CCD sensor, the present embodiment employs a
shade or hood surrounding a region between the scintillation screen
and the CCD sensor. A CCD controller then processes the electronic
representation generated by the CCD sensor, and outputs
corresponding image data.
[0007] A dual photon x-ray source is used to allow the examination
to be performed with x-rays at two different energy levels. This
source can be an x-ray tube, or a radionuclide source with a filter
element to remove one of the energy levels when desired.
Correlation of the image data retrieved using each of the two x-ray
energy levels provides quantitative bone density information.
[0008] A focusing element between the scintillation screen and the
CCD sensor can take the form of a lens or a fiber optic reducer. An
image intensifier can be used in conjunction with the CCD sensor.
The image intensifier can be a "proximity type" image diode or a
microchannel based device. It can also be directly attached to the
CCD. An image store used with the CCD controller allows
manipulation of the CCD sensor output signals by a data processor.
This includes the correlation of measurements utilizing x-ray beams
of two different energy levels. The system can also be adapted to
operate at higher shutter speeds enabling the counting of x-ray
transmissions. This provides energy measurements of x-ray
transmissions that are useful in certain applications.
[0009] In an alternative embodiment, a detector made of amorphous
silicon is used to receive and detect the radiation from the
scintillation screen to generate the electronic representation of
the spatial intensity pattern of the x-ray pattern. The amorphous
silicon detector can replace the CCD detector or it can be used to
receive the x-rays directly.
[0010] In another preferred embodiment, the apparatus of the
invention includes two scintillation screens, each of which is
associated with its own respective CCD detector or amorphous
silicon detector. One of the scintillators is reactive to
high-energy x-rays and generates an optical image of the spatial
intensity pattern of the high-energy x-ray pattern. Its associated
detector detects the image and generates an electronic
representation of the high-energy x-ray pattern. The other
scintillator is reactive to low-energy x-rays simultaneously
generate an optical image of the low-energy pattern. Its associated
detector generates an electronic representation of the low-energy
x-ray pattern. The data processor performs the correlations of the
measurements for the x-rays at two different energy levels.
[0011] An additional preferred embodiment is directed to systems
and methods of imaging spectroscopy where charge coupled device
(CCD) is optically coupled to a scintillator and measures or counts
the spatial intensity distribution of a radionuclide that has been
introduced into bodily tissue, either in vivo or in vitro. CCD's of
sufficient thickness can be used to measure gamma ray events
without the use of a scintillator in certain applications. The CCD
has sufficient resolution and sensitivity to measure such
distributions accurately, usually in less than two minutes.
Radiation sources that emit radiation having an energy in a range
between 10 and 2,000 keV, and preferably in the range between 20
and 600 keV, are delivered to the cancerous tissue or any other
suitable pathologic abnormality.
[0012] The CCD acquires "frames" of information by counting the
number of gamma-ray events over a selected period of time. Each
frame, or a sequence of frames that have been added or summed to
provide an image, can be filtered using pulse height analysis
techniques to substantially reduce or eliminate scattered
radiation. Pulse height analysis can also be utilized to
discriminate between signals having different energy levels that
contain diagnostically significant information. The system's
discrimination and energy measuring capabilities render it suitable
for diverse applications.
BRIEF DESCRIPTION OF THE DRAWINGS
[0013] FIG. 1 is a perspective view of the imaging system of the
present invention.
[0014] FIG. 2 illustrates in schematic view a bone density
measuring apparatus using a lens to focus image data from a
scintillation onto a CCD sensor.
[0015] FIG. 3 illustrates in schematic view a bone density
measuring apparatus using a fiber optic reducer to deliver an image
from a scintillation screen to a CCD sensor.
[0016] FIG. 4 illustrates another preferred embodiment for the
scintillation screen employing a fiber optic plate.
[0017] FIG. 5 is an illustration of the pixel array of a binnable
CCD sensor.
[0018] FIG. 6 is an alternative preferred embodiment to the bone
density measuring apparatus of FIG. 2.
[0019] FIG. 7 is another alternative preferred embodiment to the
bone density measuring apparatus of FIG. 2.
[0020] FIG. 8 is a perspective view of a scanning system of the
present invention.
[0021] FIG. 9 is a schematic sectional view illustrating the sensor
control system.
[0022] FIG. 10 is a schematic sectional view illustrating a frame
transfer CCD used for both emission and transmission studies.
[0023] FIG. 11 is a schematic sectional view of a CCD imaging
system for both emission and transmission studies.
[0024] FIG. 12 illustrating a process flow sequence that is used in
performing the imaging methods of the present invention.
[0025] FIG. 13 is an alternate embodiment of a CCD imaging system
that can be employed for both emission and transmission
studies.
[0026] FIGS. 14A and 14B illustrate a process flow sequence for
conducting emission and transmission studies of tissue
[0027] FIG. 15 is a schematic diagram of an alternative preferred
embodiment to the bone densitometry measuring apparatus of FIG. 2
using dual scintillation screens and dual detectors.
[0028] FIG. 16 is a schematic diagram of an alternative preferred
embodiment to the bone densitometry measuring apparatus of FIG.
15.
[0029] FIG. 17 is a schematic diagram of a variation of the bone
densitometry measuring apparatus of FIG. 16.
[0030] FIG. 18 is a schematic diagram of another alternative
preferred embodiment to the bone densitometry measuring apparatus
of FIG. 15 having dual amorphous silicon image sensors.
[0031] FIG. 19 is a schematic diagram of an alternative detection
structure including dual amorphous silicon image sensors which can
be used with the various embodiments of the bone densitometry
measuring apparatus of the invention.
[0032] FIG. 20 is another preferred embodiment of a bone
densitometer for static, scanning or stepped imaging
procedures.
[0033] FIGS. 21A-21C illustrate alternate embodiments for the
detector assembly of FIG. 20.
[0034] FIGS. 22A-22B illustrate scanning or stepped imaging
procedures.
[0035] FIG. 23 is another preferred embodiment of an imaging system
in accordance with the invention.
[0036] FIG. 24 is another preferred embodiment of an x-ray imaging
system in accordance with the invention.
[0037] FIG. 25 is a schematic diagram of a detection structure used
for bone densitometry measurements and tissue lesion imaging in
accordance with the present invention.
[0038] FIG. 26 is a preferred embodiment of the invention in which
a dual spaced array is used for digital mammographic imaging and
quantitative analysis.
[0039] FIG. 27 illustrates another preferred embodiment in which
the imaging elements in each linear array positioned at a different
angle relative to the patient and the x-ray source.
[0040] FIG. 28 illustrates a preferred embodiment in which a large
number of imaging elements are arranged in a large array conforming
in size to a standard x-ray film cassette which can have three or
more spaced linear arrays.
[0041] FIG. 29 illustrates a cross-sectional view of a linear array
using a common fiber-optic plate and scintillator.
[0042] FIG. 30 illustrates a system for translating the array
relative to a radiation source.
[0043] FIGS. 31A and 31B illustrate the process of a two step
imaging sequence.
[0044] FIG. 32 illustrates a system for sequential imaging or
scanning of tissue in which the array is moved relative to a
source.
DETAILED DESCRIPTION OF THE INVENTION
[0045] In FIG. 1 a preferred embodiment of the invention for
performing bone densitometry studies uses a detector 10 and either
an x-ray tube 12 or a radionuclide radiation source such as
Gadolinium-153. The detector 10 comprises a scintillating plate 20
which is optically coupled to a two-dimensional charge-coupled
device 24 (CCD). The CCD is a two dimensional array of detectors
integrated into a single compact electronic chip. The optical
coupling between the scintillating plate 20 and the CCD 24 is
accomplished by an optical grade lens 25. Such a lens should have a
low f-number (0.6-1.8) for adequate light collection from the
screen. The collection efficiency (E) of light from the
scintillating plate emitted in the direction of the CCD can be
computed by the equation: 1 E = tm 2 4 f 2 ( m + 1 ) 2
[0046] where:
[0047] t: Transmission factor of light through the lens
[0048] m: magnification from the Scintillating plate to the CCD
[0049] f: f-number of the lens
[0050] In an alternate approach, the optical coupling between the
scintillating plate and the CCD can be performed with a fiber optic
reducer.
[0051] Referring to FIG. 2, a bone densitometry apparatus 10 has an
x-ray tube 12 which delivers a beam of x-rays 14 towards the body
of a subject 16 being examined. The x-ray tube is capable of
emitting x-ray radiation at each of two distinct energy levels. The
two energy levels are used to obtain two distinct x-ray images of
the patient, as is discussed later. Note in comparison to FIG. 1,
the source can be positioned above the patient and the detector
below the table.
[0052] When the subject 16 is irradiated with the x-ray energy, a
percentage of the x-rays reaching the subject 16 is absorbed by the
subject's body, the amount of absorption depending on the density
of bone or tissue upon which the x-rays are incident. Since x-rays
generally travel in a straight line, the x-ray energy exiting the
subject's body on the side of the body away from the source 12 is a
spatial representation of absorption in the subject's body, and
therefore of relative tissue and skeletal densities.
[0053] To receive the x-rays passing through the subject's body, a
scintillation screen 20 is provided on the side of the patient away
from the x-ray source 12. The scintillation screen 20 is a
fluorescent material sensitive to x-rays, and when it receives
x-ray energy it reradiates visible light. The spatial intensity
patterns of the radiation emitted from the scintillation screen is
proportional to the spatial intensity pattern of the x-ray
radiation received by the screen 20. Thus the scintillation screen
20 provides an image in the visible spectrum, or alternating in the
ultraviolet or near infrared, which is regionally proportional to
the x-ray image reaching the scintillation screen 20.
[0054] A lens 22 is positioned between the scintillation screen 20
and a CCD sensor 24. The CCD sensor 24 is an array of
photosensitive pixels using closely spaced MOS diodes which convert
photons to electrons and thereby generate a discrete electronic
representation of a received optical image. The lens 22 faces the
scintillation screen and focuses the visible light emitted from the
scintillation screen 20 through the lens 22 and onto the surface of
the CCD sensor 24. In order to prevent ambient light from reaching
the CCD sensor, a shade surrounding the region between the
scintillation screen 20 and the lens 22 is provided in the form of
a photographic bellows 26. The shading of bellows 26 serves to
reduce the optical noise level of the image signal reaching the CCD
sensor 24.
[0055] Although the scintillation screen 20 absorbs most of the
x-rays incident upon it, some may still be transmitted through the
screen 20 and interfere with the optical image signal of the
scintillation screen 20. The direct interaction of x-rays with a
CCD sensor produces very bright pixels resulting in a "snow" effect
in an optical image detected by the sensor. In addition, prolonged
direct x-ray irradiation of a CCD sensor can increase its dark
current. For these reasons, an optical grade lead-glass or lead
acrylic filter 28 is positioned between the scintillation screen 20
and the lens 22 or alternatively, between the lens and the CCD. The
lead-glass filter 20 absorbs most of the stray x-rays and prevents
them from reaching the CCD sensor 24. An anti-scatter grid 29 is
used between the patient and scintillation screen for preventing
scattered x-rays from reaching the screen.
[0056] During a typical examination, the subject 16 is placed
between the x-ray source 12 and the scintillation screen 20. The
x-ray source is then activated for a short time interval, typically
one to five seconds. As x-rays are differentially transmitted and
absorbed through the body of the subject 16, they interact with the
scintillation screen 20. Upon interaction, the screen 20 emits
light in the visible part of the electromagnetic spectrum. In the
present embodiment, the scintillation screen is a terbium-activated
material and emits light in the region of 540 nm.
[0057] The light emitted from the scintillator is transported to
the CCD sensor via the lens 22. Upon interaction with the CCD
sensor 24, light energy is converted into electrons which are
stored in each pixel of the CCD sensor 24. The CCD sensor 24 of the
present embodiment consists of 512.times.512 pixels, but such
sensors come in a number of different sizes. The CCD sensor
"integrates" the image signal from the scintillation screen in that
it senses the optical image and stores charge during the entire
x-ray exposure interval. After termination of the x-ray exposure,
the discrete representation in the CCD 24 is read out by CCD
controller 30. The CCD controller 30 reads the image representation
from the CCD sensor 24 pixel and organizes it into a digital array.
The digital array, representing spatial position and x-ray
intensity, is then output to a memory or image store 32. From the
image store 32, the image can be accessed by a data processor 34
for performing image processing techniques. A cathode ray tube
(CRT) 36 is also provided to allow the image to be displayed before
or after processing by data processor 34.
[0058] Unlike other conventional detection schemes, such as film
screen radiography, CCD-based imaging provides a linear
quantitative relationship between the transmitted x-ray intensity
and the charge generated in each pixel of the CCD. After the first
high energy x-ray exposure is acquired, the resulting image is
stored in image store 32 and a second exposure with a low energy
x-ray beam is acquired with the subject 16 in the same position.
During this exposure, a low energy x-ray beam is used which is
typically at about 70 kVp with a tube current at about 1 mA. The
tube is capable of accelerating electrons at 40 kVp and up to
approximately 140 kVp. Note that the tube potential and current are
controlled by the computer menu. The low energy x-ray image is then
stored in image store 32 with the high energy exposure. Each image
provides quantitative information about the relative transmission
of x-rays through soft tissue and bone.
[0059] Once both images are obtained, comparative processing
techniques of dual photon absorptiometry are applied to determine
quantitative density measurements of those body regions scanned by
the x-rays. The correlation of two images generated by x-rays of
two different energy levels over a short time interval results in
the substantial reduction in the likelihood of systematic
pixel-by-pixel errors caused by instability of the x-ray tube
output.
[0060] Because the present embodiment of the invention is concerned
with an area detector as opposed to a scanning detector, the
measurement time necessary for a densitometry examination is
greatly reduced. Rather than scanning across the region to be
examined in a rectilinear fashion, the entire region is irradiated
simultaneously and the resulting image processed simultaneously.
Typically, the entire procedure using the present dual photon
technique lasts 30 to 60 seconds, depending on the power of the
x-ray tube and processing speed of the supporting electronics.
[0061] FIG. 3 shows an alternative embodiment to that of FIG. 2. In
this embodiment, the x-ray tube source 12 of FIG. 2 is replaced
with radionuclide source 40. The radionuclide source is
gadolinium-153. Gadolinium-153 emits photons simultaneously in two
energy bands, a lower energy band of 44 keV and an upper energy
band of 100 keV. Thus, the gadolinium source is a dual photon
radiation source. In order to allow the images from the two
different energy levels to be obtained separately, an x-ray filter
42 is placed between the source 40 and the subject 16. In the
present embodiment, the filter 42 is copper or a K-edge filter, and
eliminates nearly all of the low energy (44 keV) emission from the
beam. Removal of the filter restores the beam to its dual energy
nature. The filter 42 is implemented as an electromagnetic shutter
which may be opened and closed in the line of the x-ray beam. A
high energy image is acquired first with the filter shutter closed,
after which an image is obtained using the dual energy beam with
the shutter open.
[0062] Both electronic images are stored, and an image
representative of the transmission of only the low energy photons
is obtained by electronically subtracting the high energy image
from the dual energy image with the data processor 34. Once both
images are obtained, comparative dual photon processing techniques
are used to make quantitative density calculations.
[0063] An additional feature of the embodiment of FIG. 3 is the
replacement of the lens 22 of the FIG. 2 embodiment with a fiber
optic reducer 44. The fiber optic reducer 44 is a focusing device
consisting of a large array of optical fibers packed tightly
together, and leading from the scintillating screen 20 to the CCD
sensor 24. Near the CCD sensor 24, many of the fibers can be fused
together, thus combining the signals present on individual fibers.
The effect is a compression of the image from the input of the
reducer 44 at the scintillation screen 20 to the reducer output at
the CCD sensor 24. In this manner, the reducer 44 effectively
focuses light from the scintillating screen 20 onto the CCD sensor
24 without the necessity of a lens for the focusing region.
[0064] Although they are shown together in FIG. 3, it is not
necessary to use the fiber optic reducer 44 with the radionuclide
source 40. Either element can be substituted into the configuration
of FIG. 2 individually. The x-ray filter 42, however, should be
used with the radionuclide source 40 to provide a dual photon
discrimination capability. Note, however, that pulse height
analysis can be performed in conjunction with the embodiment of
FIGS. 10 & 11.
[0065] FIG. 4 shows an alternative to the scintillation screen 20
of FIGS. 2 and 3. The screen 48 depicted by FIG. 3 is a
scintillating fiber optic plate. The plate 48 is a fiber optic
faceplate consisting of scintillating fibers 50 running through the
plate. The fiber optic plate is optically interfaced to the CCD in
essentially the same way as the scintillation screen 20 of FIG. 2,
but the fiber optic plate 48 allows for greater quantum efficiency
due to increased x-ray stopping capability.
[0066] Shown in FIG. 5 is a representation of the pixel array of
the CCD sensor 24. The array shown in FIG. 5 is only 10.times.10
for illustrative purposes, and the actual array can be of different
dimension. Each pixel in the array is an individual photosensitive
element which contributes to the overall image detected by the
array. A feature of the CCD sensor of the present embodiment is the
capability of the pixels of the sensor 24 to be "binned" together.
The binning of the pixel array refers to the ability of the sensor
electronics to combine groups of pixels together to form "super
pixels" which are then identified as single picture elements.
[0067] Charge is binned by combining charge packets contained in
two or more adjacent potential wells into a single potential well
during charge readout. Serial and parallel binning can be combined
to perform two dimensional binning from any rectangular group of
wells or detector elements.
[0068] The dark lines in the binnable array of FIG. 5 illustrate
where individual pixels might be grouped together. For example, the
four upper left hand corner pixels 50 can be binned together
through control of the CCD sensor 24 to form a super pixel. The
super pixel is then identified by the CCD electronics as a single
pixel, the light intensity reaching each pixel 50 being averaged
across the surface of the entire super pixel. In this manner, the
dimension of the array can be electronically controlled. As can be
seen in FIG. 5, if groups of four pixels are binned together across
the 10.times.10 array, the overall array dimension becomes
5.times.5. Although the binning of the CCD sensor 24 reduces the
resolution of the pixel array, the relative percentage of noise is
also reduced, thus providing an improved signal to noise ratio.
[0069] The following x-ray data acquisition approach is an
alternative to the one described above. In this approach, an image
is acquired at high energy and the CCD is read in the normal
non-binned mode. Due to the high penetration of the high energy
beam through the body, the x-ray fluence exiting the body is high
as compared to that of the low energy beam. Therefore, the
resulting charge signal per CCD pixel is relatively strong. This
image is stored as the high energy image. Also, this image is used
in order to compute the area of the bone to be measured by manual
selection of the region of interest or by automatic edge detection.
Therefore, we take advantage of the high resolution image for
greater accuracy in the measured bone area. Previously, the
accuracy and precision of bone density measurements are limited to
a great extent by suboptimal spatial resolution. The next image
which is acquired with low energy is read out by the pixel binning
approach, e.g., using a 2.times.2 pixel binning. The transmission
of the low energy beam through the body is low as compared to the
high energy beam. Therefore, in order to record a strong signal in
each CCD pixel we must increase the radiation dose.
[0070] Alternatively, the binning technique can be used for the low
energy in order to increase the signal to noise ratio and to
decrease the radiation dose. This dual mode acquisition procedure
is a very powerful tool for improving the signal to noise ratio and
lowering the radiation dose to the patient.
[0071] Although the arrangement of optical elements as shown in
FIGS. 2 and 3 represent preferred embodiments, the functionality of
the system is not dependent upon such an in-line type of optical
transmission. FIG. 6 shows an alternative arrangement of optical
elements where the CCD sensor 24 is set at an angle relative to
scintillation screen 20, and mirror 52 is used to reflect the
radiation given off by the scintillation screen toward the CCD
sensor 24. Lens 22 is shown between CCD sensor 24 and mirror 52 and
focuses the image onto the CCD sensor. However, the focusing of the
scintillation screen image can take place before or after the image
reaches mirror 52. In fact, the mirror itself may be shaped to
provide focusing of the image from the scintillation screen 20.
[0072] FIG. 7 shows another alternative arrangement of optical
components. In FIG. 7 the subject 16 is suspended by a support 54
which is transparent to x-rays. The support 54 keeps the subject 16
elevated a distance above scintillation screen 20. As the x-rays
reach scintillation screen 20, the screen 20 reradiates image data
from the same surface upon which the x-ray radiation is incident.
Mirror 52 is now aligned to reflect this image towards CCD sensor
24 which collects the image as focused through lens 22 to be
processed by the CCD controller 30.
[0073] As with the arrangement of FIG. 6, the focusing of the image
from the scintillation screen 20 may take place before or after it
is reflected by the mirror 52, or may be focused by the mirror 52
itself. In addition, any of the optional elements previously
discussed may be substituted into the arrangement of FIG. 5 or FIG.
7. This includes the x-ray absorbing screen 28, the anti-scatter
grid, the fiber optic reducer 44, and the fiber optic faceplate
48.
[0074] A very effective, radiation dose-efficient approach for
reducing x-ray scatter and increasing the dynamic range of
electronically acquired x-ray images is the use of a slit-scan
method. In this approach, a fan beam of x-rays is scanned over the
patient and a linear array of detectors is used to detect the
transmitted radiation. In typical applications the length of the
detector restricts the width of the area that can be covered with
one pass. Also, many small linear CCD or photodiode arrays are used
to form a line of detection. This results in a rather complex
detector assembly. If cooling of the detector assembly is required,
it is difficult to accomplish for such an extended detector. Also,
image intensification by using an electronic intensifier becomes
difficult and very costly.
[0075] An alternative embodiment for dual energy bone densitometry
takes advantage of the merits of slit-scan geometry without using a
linear CCD or photodiode array. This approach is illustrated
schematically in FIG. 8. An area CCD sensor 64 is used in
conjunction with a line-to-area fiber optic converter 62. This
converter can be made of flexible or rigid optical fibers with
cladding of lower index of refraction than the core material. As
shown in FIG. 8, the CCD 64 is divided into a number of rows and a
fiber optic ribbon is optically coupled or bonded to each row. The
coupling of the CCD 64 to the converter 62 can be accomplished
using the various systems described in connection with other
embodiments. An extramural absorber can be used to prevent light
crossing from one fiber to another. The other ends of each ribbon
are arranged in tandem to form a linear sensor. In front of the
linear sensor (input end), an x-ray converting scintillator 60 is
used such as gadolinium oxysulfide activated with terbium (GOS:Tb).
Alternatively, a scintillating fiber optic plate can be used for
improved quantum efficiency at higher energies. A linear x-ray
sensor with a very compact area detector is employed with the
slit-scan embodiment.
[0076] A typical linear detector of this type comprises a few
ribbons in tandem along the length of the slit, and from one to a
multitude of ribbons across the width of the detector slit.
[0077] In a typical example, consider a 512.times.512 pixel CCD
where each pixel has an area of 20.times.20 microns. A fiber optic
bundle with individual fibers of 60 microns in diameter is used for
the embodiment. On the CCD each fiber will cover an area of
approximately 3.times.3 pixels. Perfect alignment between each
pixel and fiber is desirable but it is not essential for this
application. Close packing of the fibers will result in an array of
170.times.170 or a total of 29,127 fibers covering the entire area
of the CCD. Each ribbon of fibers corresponds to one row consisting
of 170 fibers and covering approximately 512.times.3 pixels on the
CCD. If all ribbons emerging from the CCD were arranged in tandem,
the linear sensor would be approximately 175 cm in length.
Alternatively, the ribbons can be arranged with a small number in
tandem and a small number across the width of the slit. Using the
above CCD, a 15.3 cm linear detector can be made with approximately
15 ribbons in tandem thus using only a small fraction of the CCD
area.
[0078] Full use of the CCD area can be made by stacking the ribbons
in groups of 15, (one ribbon per CCD row), thus creating a
quasilinear detector consisting of an array of 2,550.times.11
fibers optically coupled to an x-ray scintillator. The dimensions
of this slit detector will be 153.times.0.66 mm with a total
sensing area of 1.0 cm.sup.2. It is important to note that the
total sensing area of the slit must be approximately equal to the
total area of the CCD and the linear dimensions of the CCD. A wider
or longer slit will result in a larger area at the output end. In
this case, a larger CCD can be used for a fiber optic reducer
optically bonded between the fiber optic converter and the CCD.
Alternatively, the converter itself can be tapered to match the
size of the CCD. For higher spatial resolution the fiber optic
converter is made with optical fibers of smaller diameter (5-6
microns).
[0079] If higher signal amplification is required for some high
detail low dose applications, a proximity focused image intensifier
can be optically bonded between the fiber optic taper and CCD or
between the fiber optic converter and fiber optic taper. The image
intensifier can be a proximity diode type or a microchannel plate
device, both commercially available. Alternatively, an integral
assembly of CCD and intensifier can be used commonly called an
"intensified CCD". Another approach is to use a lens coupling
between the output surface of the fiber optic converter and the
intensified or non-intensified CCD.
[0080] Cooling of the CCD can be accomplished easily by a
thermoelectric cooler. Cooling is required only when very high
contrast resolution is required and the image acquisition time is
relatively long. If the CCD is read out at 500 kHz
(5.times.10.sup.5 pixels/sec), an area of 150 mm.times.150 mm of
the subject can be scanned in approximately 114 seconds
(approximately 2 minutes). Faster scanning is attainable by
increasing the readout rate of the CCD.
[0081] Alternatively, a frame transfer CCD such as the one
illustrated in FIG. 10 can be used for faster scanning. This device
uses one half of its sensing area for storage and not for sensing.
In this way the transfer of the image from the sensing area 91 to
the storage area 93 is accomplished in a few milliseconds. A
smaller CCD such as a 128.times.128 or a 64.times.64 element could
be used for this purpose in a similar arrangement as with the
512.times.512 CCD. Also, larger area CCDs can be used for this
purpose. Pixel binning as described previously can be applied in
this detection approach. A Gadolinium-153 (Gd-153) radiation source
can be used as described in previous sections in place of an x-ray
tube. The Gd-153 source is a small pellet or a collimated line
source parallel with the long dimension of the detector.
[0082] The line to area conversion design enables us to remove the
CCD from the direct path of the x-ray beam, thus it allows for easy
shielding of the CCD from direct x-ray interactions. This prolongs
the useful life of the CCD and it alleviates the "snow" effect
which results from direct interactions of x-rays with the sensor.
Moreover, this approach allows for greater light transport
efficiency between the scintillator and CCD than lenses or fiber
optic tapers. Note that the pixel binning approach enables the
operator to select the desired spatial resolution and contrast
without any mechanical modifications on either the x-ray beam or
the detector collimator. The pixel size of the detector which
determines resolution and contrast can be controlled by a command
from the computer. This x-ray imaging modality can be used very
effectively to optimize the scan depending on the patient size, and
medical history.
[0083] An alternate approach provides an improved rectilinear
scanning method for quantitative x-ray radiography. In this
embodiment, a two dimensional CCD optically coupled to a
scintillator is used as the detector of x-rays in a rectilinear
scanning mode. The CCD may be a full frame or a frame transfer
device. The frame transfer CCD will enable faster data scanning and
acquisition.
[0084] The CCD scintillator assembly is extremely critical to the
performance of the system. Direct optical bonding of a
polycrystalline scintillator such as gadolinium oxysulfide with the
CCD is possible but this approach is not efficient in shielding the
CCD from direct x-ray interactions. If the thickness of the layer
is increased the spatial resolution of the x-ray images degrades
due to light diffusion. The use of a scintillating fiber optic
plate between the polycrystalline scintillator and the CCD provides
a solution to this problem.
[0085] A scintillating fiber optic plate is a fiber optic faceplate
designed to convert x-rays or U.V. light into green light with peak
emission at about 550 nm. This faceplate is manufactured with the
extra mural absorber to prevent light diffusion between individual
fibers. The area of the scintillating fiber optic plate must cover
the CCD completely. The desirable thickness depends on the energy
of the x-ray radiation. A thickness of 5 to 10 mm is preferable but
a thinner or thicker plate can be used. The use of a very thick
scintillating fiber optic plate such as 10 mm or 20 mm will
eliminate virtually any undesirable direct x-ray interactions with
the CCD. The scintillating fiber optic plate can also be used
without the thin layer phosphor. However, the combination of the
two will produce better image quality at a reduced radiation dose
to the patient. Alternatively a conventional fiber optic plate can
be used as a substrate to the scintillating fiber optic plate. The
optical coupling of the polycrystalline phosphor on the fiber optic
can be accomplished by direct deposition techniques or by using an
optical adhesive.
[0086] In an alternate approach, a bent fiber optic bundle can be
used between the scintillator and the CCD. The geometry of the bent
bundle allows for extremely effective shielding of the CCD from
extraneous x-ray radiation. A lens coupling between the CCD and the
fiber optic converter can also be used. For improved sensitivity, a
proximity focused image intensifier, an image diode or microchannel
plate can be used at the input end of the fiber optic or between
the fiber optic bundle and the CCD. A preferred approach is to use
the intensifier at the input end. A scintillator can be optically
bonded to the input of the intensifier or an intensifier with a
scintillating fiber optic input plate can be used.
[0087] The x-ray tube is aligned in a C-arm configuration with the
detector. The x-ray beam is approximately congruent with the area
of the detector which is approximately 1.times.1 cm at the detector
plane. As x-rays are transmitted through the patient, some
(20%-60%) are absorbed by the primary polycrystalline scintillator
producing visible light. This light is transmitted through the
optically transparent fiber optic faceplate in the direction of the
CCD. The x-rays not interacting with the primary scintillator will
be absorbed by the fiber optic faceplate. If a scintillating fiber
optic faceplate is used, these x-rays will be absorbed in the
fibers thus producing additional scintillations. Therefore, the
scintillating fiber optic plate acts as a light conduction device,
x-ray shield, secondary x-ray detector and an x-ray signal
amplifier.
[0088] Upon interaction of the x-ray induced light with the
photosensitive surface of the CCD an electron charge is generated
which is proportional to the number of x-ray interactions in the
scintillators. The cumulated charge on the CCD is then read out.
However, in this rectilinear scanning mode, each CCD readout will
correspond to a small segment of the total image, approximately one
square centimeter. Therefore, the entire image is acquired by
spatial additional of each image segment. For example, if a
15.times.15 cm field is covered and the sensor area is
1.0.times.1.0 cm, 15.sup.2 (225) segments must be acquired and
synthesized. A 512.times.512 pixel CCD operating at 500 kHz will
read out each segment in 0.5 seconds and will require about 2
minutes for the entire scan at a scan speed of about 2 cm/sec.
Faster scanning is attainable by increasing both the scanning speed
and the readout rate of the CCD.
[0089] A dual-energy scan will be acquired by first scanning the
entire area at high tube potential, typically 130 kVp without
binning and then repeating the scan at low tube potential,
typically at about 70 kVp with binning. An automatic slide
mechanism places high aluminum filtration for the high energy beam
and less filtration for the low energy beam as described
previously. The images of each energy level are stored in the
computer for subsequent dual photon analysis. Pixel binned
acquisition will be possible at both energies for improved
precision. Where both high and low energy images are identically
binned, this produces an exact correlation between the images
produced. A third high energy-high resolution image can then be
used to define the outline of the object being scanned. Note that a
gadolinium isotope source with a shutter can be used.
[0090] Alternatively, the energy level of the tube can be switched
from low to high for each segment of the acquisition and each
segment representing high and low energy is stored for subsequent
analysis.
[0091] An alternate approach employs light intensification from the
screen to the CCD sensor. In this approach, an electrostatically
focused image intensifier (In FIG. 2) is employed as the primary
detector in place of the scintillating plate. This intensifier
preferably employs Cesium iodide input phosphor with an approximate
diameter of 15 cm and thickness of 0.3-0.5 min. The high voltage of
the image intensifier tube can be reduced to approximately half the
normal value. A reduction in the image intensifier accelerating
potential will contribute to an improvement in the image contrast
characteristics and dynamic range of the device. The CCD sensor is
optically coupled to the output phosphor of the image intensifier
by a fast lens with an f-number of about 1:1.0. Due to the high
signal intensification, cooling of the CCD is not essential but it
can be applied if very low thermal noise levels are desirable. The
use of an intensifier allows for the use of a CCD with lower noise
performance characteristics, thus lowering the cost and complexity
of the instrument.
[0092] Ideally, the detected signal is produced by x-rays that have
been transmitted through the body without any scatter interaction.
Detection of large amount of scatter events will result in
non-linearities and in a reduction in the dynamic range. Effective
suppression of scatter is accomplished by using a small field of
view, typically 10 cm.times.10 cm and by using a air gap
(approximately 20 cm) between the patient and the scintillating
plate. Alternatively a small field of view can be used in
conjunction with a linear or crossed antiscatter grid.
[0093] An internal instrument stability control system has been
incorporated to provide a means of automatic compensation for any
instabilities in the x-ray tube potential and current. The
stability control device is not essential for the operation of any
of the described techniques but it provides better reliability and
precision in the measurement of bone density. A schematic
represen-tation of the proposed device is shown in FIG. 9. The
output of the x-ray tube 12 is monitored by a pair of x-ray sensors
70 placed at a secondary x-ray beam port 78 adjacent to the main
beam port 80 near the tube window. The sensors can be silicon
diodes, cadmium telluride radiation sensors or any other solid
state x-ray sensor. Alternatively, a pair of compact
photomultiplier-scintillators or a photodiode scintillator assembly
could be used. Both detectors operate in the charge integration
mode and the detected signal is continuously monitored as a
function of time during the entire scan for each energy. This time
varying signal is digitized and stored in the computer memory. The
change in the filtration of the secondary beam with energy is
identical to that in the main beam because it is controlled by the
same filter changing mechanism. As described further in connection
with FIG. 12 the sensor system can be used to normalize the
detected information or to control operation of the x-ray source to
prevent or reduce unwanted variations in the source output.
[0094] In front of one of the sensors 70 an amount of polymethyl
methacrylate 86 is placed to simulate an average thickness of soft
tissue. In front of other sensor 70 an amount of bone simulating
material 84 is placed in an amount equivalent to that encountered
in the spine or femur. Various hydroxyapatite-epoxy mixtures are
commercially available for bone simulation in x-ray imaging.
Therefore, a secondary detection system with a bone standard of
known density and a soft tissue equivalent thickness is provided in
this embodiment.
[0095] The signals from each sensor 70 can be used to compute the
density of the bone internal standard as a function of time during
the scan. Any deviations from a constant density of this standard
are due to changes in either the energy or intensity of the x-ray
emission. Each value of bone density computed in the patient scan
corresponds to a computed value of the bone standard. Therefore,
each computation of bone density derived from a pair of high and
low energy CCD frame acquisitions can be corrected or normalized by
using the deviation from the density of the internal standard. For
example, if the value of the bone standard during the rectilinear
scan deviated by plus 3% in a given area of the image, the computed
bone density of the patient scan must be corrected by that amount
in this area. This internal reference approach can be used with all
stationary and scanning embodiments described herein.
[0096] In conjunction with the above calibration approach, a number
of strips 72 (square rods) of bone simulating epoxy material, or
aluminum of equivalent x-ray absorption are placed under the table
73 which run in the direction of the scan for the slit scan
approach. Each linear strip has a different thickness or bone
equivalent density. As the x-ray tube and detector assembly is
scanned over the area to be tested, each set of rods are scanned
and their density computed. The consistency of the measured
densities of these rods is used to ensure proper operation of the
system. This set of standards can be placed anywhere from the x-ray
exit port 80 to the edges 74 of the detector 76.
[0097] The imaging of radionuclide distributions in biological
tissues or specimens is a routine task performed in virtually all
biomedical research laboratories by the well established technique
of autoradiography. In this procedure, a thin slice of a specimen
is placed in contact with photographic film thus allowing the
radiation from the specimen to expose the film. Subsequently, the
film is processed by standard chemical development techniques,
manually, or by using an automatic processor. Frequently, an
intensifying screen is used in order to enhance the absorption
efficiency of the image receptor and for a reduction in exposure
time. Intensifying screens are especially useful when images of
relatively high-energy gamma or x-ray emissions are recorded
(20-200) keV. Also they can be useful for high energy
electrons.
[0098] Autoradiography produces images reflecting the
biodistribution of a radionuclide and it has been established as a
powerful tool in many biomedical disciplines. Its major
shortcomings relate to problems with quantization of the relative
or absolute concentration of radionuclide in an area of interest.
This difficulty arises from the non-linearity of photographic film
typically used and in reciprocity law failure when intensifying
screens are used. Moreover, the development temperature, and in
general, the condition of the processing chemicals have an
influence on the film fog level and contrast. All these factors
render quantization a very difficult and time consuming task which
becomes vulnerable to many uncertainties in quantitative
autoradiography. Despite this problem, several investigators have
digitized film autoradiographs by using microdensitometers or video
cameras for both quantization and image enhancement.
[0099] In autoradiography, the image represents areas where the
radiotracer has been extracted. The anatomical information on the
original tissue slide is not transferred with great detail in the
autoradiograph. For proper interpretation, it is necessary to
observe the tissue slide and autoradiograph side by side in order
to correlate radiotracer distribution with anatomy. Often it is
necessary to superimpose the slide with the autoradiograph in order
to identify the exact anatomic location of the radiotracer. In this
process the accuracy in assigning an anatomic location to the
tracer is severely compromised.
[0100] One of the most important problems with autoradiography is
the long period of time required in order to expose the film. In
most applications this time ranges from a few hours to several
days, even weeks in some cases. Therefore, the technician may have
to wait for a few days in order to find out whether an exposure has
to be repeated.
[0101] Autoradiography does not relate to in vivo imaging of
radionuclide distributions in humans or animals. Rather it relates
to detecting radioactive distributions in excised samples. All
available film-screen image receptors have extremely low quantum
efficiencies for most gamma emitters commonly used for this
purpose. Moreover, the presence of a large volume of tissue results
in enormous amount of gamma ray scatter which will reach the image
receptor and degrade the contrast and spatial resolution. The
film-screen receptors do not have energy discrimination
capabilities, therefore scattered events cannot be rejected. The
use of a collimator to suppress scatter will result in a dramatic
reduction in geometric efficiency.
[0102] Thus the present invention, in its various embodiments,
provides an effective means for performing autoradiography by
providing a compact device that performs the data acquisition for
autoradiography quickly and can superimpose both emission and
transmission studies to correlate the emission image with the
anatomical features of the object under examination. The
embodiments described in connection with FIGS. 10 and 11 below can
be used to perform autoradiography procedures.
[0103] Radionuclide imaging of humans and animals is performed on a
routine basis by using the Anger camera, most commonly referred to
as a "Gamma Camera". The gamma camera has a quantum efficiency in
excess of 50% for the most commonly used radionuclides and it has
the capability of discriminating scatter from primary photons by
pulse-height analysis of each detected photon. The intrinsic
spatial resolution of the gamma camera is approximately 3.5 mm. The
total spatial resolution of the camera, including the degradation
due to its collimator, can vary from 5 mm to 12 min. Modem gamma
cameras can detect photons at the rate of 25,000 counts per second
(cps) without significant dead time losses. At higher count rates,
significant deviations are observed between true and detected
events. This is due to limitations inherent in the design of both
the detector assembly and processing electronics.
[0104] The following presents a further embodiment relating to
imaging of radionuclide distributions in tissue samples and in vivo
quantitative imaging of humans and animals. This procedure employs
a charge-coupled device to detect and process information to
provide, in essence, a compact "gamma camera" using a highly
sensitive stationary (or scanning) detector to conduct both
emission and transmission studies at count rates up to 10.sup.6 of
the object being examined.
[0105] Existing gamma cameras have limited spatial resolution,
limited capability to perform in high count-rate conditions and it
cannot be used to record x-ray transmission (radiographic) images
with any degree of acceptable detail to satisfy radiographic
imaging standards. Therefore, the recording of a high quality
radionuclide (physiologic) image and a radiographic (anatomic)
image with the same detector for accurate correlation of the
physiologic and anatomic image remain difficult. Where very high
detail is necessary, the gamma camera is generally not capable of
producing better than 5 mm resolution even under the most favorable
conditions. Therefore, the imaging of small parts of the body or
imaging small animals like mice cannot be performed with any
reasonable detail using the gamma camera. This also applies for the
imaging of tissues containing radioactive materials.
[0106] The following procedures enable the acquisition of high
detail radionuclide images and the option of combining them with
the x-ray radiographic images with the same detector. This approach
employs a novel acquisition scheme that enables imaging
spectroscopy of gamma rays, x-rays or nuclear particles by using a
CCD. CCDs have been employed in the past without a scintillator for
imaging spectroscopy of very soft x-rays, up to the energy levels
of about 6-9 keV. However, above this energy, the CCD becomes
virtually transparent to x-rays or gamma-rays. Generally,
scintillators have not been used in conjunction with a CCD for
imaging spectroscopy because it is believed that the conversion
from gamma-rays to light will destroy the useful information
carried by the interacting gamma-ray or x-ray. Therefore, imaging
spectroscopy of gamma-rays or x-rays in the energy range of about
10 keV to 2,000 keV with a CCD has not been explored. Also,
alternating the mode of operation from a counting, energy sensing
detector to an integrating detector for radionuclide and
radiography, respectively, presents a useful procedure for imaging
spectroscopy. Note, however, that the counting procedure can also
be used in certain x-ray transmission measurements to measure the
energy thereof.
[0107] When light interacts with the sensitive surface of the CCD,
it generates a charge which remains stored in the pixel where this
interaction occurred. As with previous embodiments the magnitude of
the charge is directly proportional to the detected intensity of
light. Each pixel is represented by its two-dimensional coordinates
and by an intensity value. The energy required to produce an
electron in the sensitive silicon surface of the CCD is about 3.65
eV.
[0108] This value enables the determination of the energy of
detected photons if the system can either detect one photon at a
time, or if the number of the photons detected per pixel is known.
This provides for imaging of radionuclide distributions with a
simultaneous measurement of the energy of the detected events. This
procedure is termed "Imaging Spectroscopy" and provides a technique
using gamma rays, beta-rays, and x-rays in conjunction with CCD
technology.
[0109] The upper energy limit of soft x-ray imaging is between 5-10
keV. At 10 keV, the quantum efficiency of a CCD is approximately 5%
and it diminishes rapidly at higher energies. The small fraction of
the total number of events interacting with the CCD will result in
a high partial energy transfer to the sensor with losses in
proportion with the energy and the signal. Therefore, when the CCD
is used as the primary detector of high energy photons or
particles, it is virtually unusable for performing imaging
spectroscopy. The following procedure provides high resolution
imaging spectroscopy using a CCD that is suitable for many
applications including position emission tomography and nuclear
particle imaging.
[0110] A schematic of the device is shown in FIG. 10. An important
component of this device is a CCD 98 with low readout noise, high
charge transfer efficiency and dark current levels. A CCD with less
than 10 electrons/pixel (RMS) readout noise is suitable for this
purpose. The dark current can be reduced to less than 0.6
electrons/sec at -40.degree. C. by a compact thermoelectric
cooler.
[0111] In one embodiment of this method, a thin scintillator 104 is
used as the primary detector of x-rays. One such scintillator can
be a layer of gadolinium oxysulfide or thallium activated cesium
iodide or any of the commonly available phosphors. The scintillator
104 is bonded to a fiber optic faceplate 106 and the faceplate is
bonded to an image intensifier 96. The intensifier is bonded to a
second faceplate 106 that is bonded to bundle 102. Optical bonding
of this type is well established. To further illustrate this
embodiment the sensitive area of the scintillator 104, faceplates
106, image intensifier 96, fiber optic coupler 102, and CCD 98
have, identical dimensions. Note that a collimator 94 can be
mounted on the lead enclosure 100 and is used during the
transmission study, and depending on its configuration, can also be
used during the emission study. Note that the collimator 94 can
optionally be removed during emission studies.
[0112] When an x-ray photon within the rays 14 interacts with the
scintillator 104, it produces light with intensity which is
proportional to the energy of the x-ray. This light is transported
through the fiber optic faceplate 106 and interacts with the CCD
98. The interaction of optical photons in each CCD pixel will
produce a number of electrons in direct proportion to the number of
optical photons and to the energy of the detected x-rays 14 or
gamma-rays 92 that are produced by the isotope that has collected
in the lesion 90. Isotopes commonly utilized include TC 99m or
I-125. The following example as a first order approximation of the
expected energy resolution from the detector.
[0113] A 60 key x-ray interacts with the scintillator resulting in
3000 optical photons. Approximately one half of these photons are
emitted in the direction of the CCD. Assuming a Lambertian
distribution of the emitted photons from the screen, the
transmission through the fiber optic plate is approximately 40%.
Therefore, 600 optical photons will be arriving at the CCD. The
quantum efficiency of the CCD is approximately 40%, therefore only
240 photons will be detected in one pixel.
[0114] It can be shown that the energy resolution can be in the
order of 10% which is approximately twice that attained with
conventional NaI-crystal spectrometers at this gamma-ray
energy.
[0115] FIG. 11 depicts an alternative embodiment in which a "pin
hole" collimator 112 with shutter 110 is used in performing an
emission study of lesion 90 or any selected organ. The emission
from the lesion or organ impacts the scintillator 104, into housing
100, through the fiber optic reducer 116, coupled to the
intensifier 118, and than directed off mirror 124, lens system 120,
and onto a cooled CCD 120.
[0116] This procedure produces radionuclide scintigraphy with
spatial resolution in the order of about 1 millimeter or less, and
transmission images with resolution in the order of 0.2
millimeters. The spatial resolution and sensitivity of the detector
will be selectable for both emission and transmission modes via
pixel binning. The detector operation will be selectable for
pulse-height analysis or integration. For x-ray transmission
imaging, the integrating mode of operation is preferred. Note that
during x-ray transmission imaging, the pin hole collimator will be
removed. Emission imaging of thick tissues requires a collimator,
either a multihole type or a pinhole collimator. Very thin
specimens can be imaged without a collimator by placing them very
close to the scintillator.
[0117] This camera has the capability of detecting very high count
rates. In conventional gamma cameras, each x-ray photon interaction
occupies the entire scintillator and electronics for a period of
time of 1 to 8 microseconds after it is detected. In the present
method, due to the multiple detectors, higher count rates can be
handled due to the multiple detectors, and higher count rates can
be handled without using a scintillator with short decay time.
Count rates up to 10.sup.6 counts per second can be acquired with
very low probability (less than 1 %) of detecting 2 gamma ray
events in one pixel when operating in the pulse-height analysis
mode.
[0118] Note the scintillator can be bonded directly on the fiber
optic bundle without the use of an image intensifier. Also, the
scintillator can be bonded directly on the CCD without the use of a
fiber optic bundle. A frame transfer CCD is a preferred approach,
but a full frame CCD can be used.
[0119] The following "shutter" methods can be used (a) a frame
transfer CCD; (b) a gated image diode, or microchannel intensifier;
or (c) a liquid crystal shutter with very thin window or fiber
optic window. The liquid crystal shutter can be positioned between
the fiber optic bundle and the scintillator.
[0120] Note that the system has applications for small animal
imaging, skeletal imaging, monitoring of fracture healing, thyroid
scintigraphy, Bremsstrahlung imaging of beta emitters within the
body (radiation synovectomy), intraoperative imaging probe,
radionuclide angiography, small parts imaging, and pediatric
nuclear imaging.
[0121] FIG. 12 illustrates in schematic form several methods that
can be used in performing quantitative imaging in accordance with
the various embodiments of the invention.
[0122] Note that one can use either a stationary source and
detector to project radiation 130, or a scanning source and
detector assembly to scan the object being examined 132.
[0123] Both stationary and scanning embodiments utilize a CCD
detector that transfers the detected information to a memory 140.
The information can be binned or processed 142 to accomplish
various tasks. This processing can include the application of
software modules to correct for non-uniformities in the source or
collection components, or to identify events where light from one
gamma-ray interaction has spread to a number of neighboring pixels.
Clusters of pixels with high intensity can be identified as primary
events and low intensity clusters can be identified as scattered
radiation and be eliminated by a filter.
[0124] Quantified information such as an intensity histogram (i.e.,
a pulse height spectrum) can be generated 146 and a display of the
object can be generated 144 with the unwanted pixels removed.
[0125] After each set of data is produced in both the stationary
and scanning embodiments, the conditions for operation can be
modified 138 to produce an image at a different energy level, to
perform an emission or transmission study, or to rotate the source
and detector assembly relative to the object under study to produce
three dimensional images or two dimensional images at different
angles.
[0126] The emission and transmission studies can be displayed alone
or superimposed. Due to the binning capability of the system a one
to one correspondence exists between both emission and transmission
images that was previously not possible. This high resolution image
can be color coded to distinguish between the emission and
transmission images.
[0127] Another preferred embodiment is illustrated in FIG. 13 where
a full frame or frame transfer cooled CCD 150 with a transparent
scintillator 152 bonded on the sensitive surface of the CCD, or to
an image intensifier 154, as shown. The scintillator 152 is
preferably emitting anywhere from the UV blue to the red regions of
the spectrum upon stimulation with x-rays or gamma-rays. The
preferred scintillator is one emitting in the green such as CsI
(TI) or Cadmium tungstate, or alternatively a gadolinium based
ceramic scintillator available from Hitachi Corporation. This
scintillator has about twice the density of sodium iodide or
CsI(TI) and has higher efficiency. A fiber optic plate (straight or
reducing) can be incorporated between the CCD and scintillator.
Alternatively, an electrostatic image intensifier 154, or image
diode intensifier, can be incorporated between the scintillator and
the fiber optic plate. The scintillator 152 can be optically
transparent plate or comprise a fiber optic array with fibers
ranging in diameter from 0.006 mm to one or more millimeters. The
thickness of the plate can be in the order of 0.5 mm to 5 mm.
[0128] Another preferred embodiment employs a CCD of the type
described above but in conjunction with an electrostatic
demagnifying image intensifier. The optical coupling of the CCD is
accomplished by a fast lens at the output end of the image
intensifier or by a fiber optic plate between the output screen and
the CCD.
[0129] The process of obtaining a desired image includes the
initiation of acquisition with the CCD for about one second or at a
desired binning configuration, typically coarser than 2.times.2
pixels. Shorter acquisition time will be required for high
count-rates and longer acquisition time is tolerated for low count
rates. The optimal acquisition time for a particular application
can be determined empirically by acquiring a few test frames and
search for coincident events within individual pixels. Very short
acquisition times (less than 1 millisecond) are easily attainable
by using a fast mechanical shutter, an electro-optical shutter, or
by gating the image intensifier tube. This enables acquisition with
spectroscopy capability even at very high count-rates. Each
acquisition "frame" will record from a few hundred to a few
thousand counts. After acquisition, each frame is stored in the
computer memory for subsequent processing. Depending on the
application, the total number of frames for a complete acquisition
can vary, for example, from ten to a few hundred.
[0130] Each gamma-ray event in a given frame stored in the computer
is represented by its x and y coordinates and by an intensity value
(z) which is the number of electrons generated in this area of the
CCD. The z value is directly proportional to the energy of the
gamma ray (or x-ray). The number of electrons generated from each
interaction should be confined to one pixel or group of binned
pixels forming a "superpixel". In a significant percentage of
interactions, the electrons generated from a single gamma-ray
interaction can be split between two or three pixels or
superpixels. These split events form clusters in the image matrix
which can be easily identified by the computer software and
assigned an x and y coordinate.
[0131] In one embodiment, as shown in the process flow sequences of
FIG. 14, pulse height analysis uses the value of these neighboring
pixels which are summed to produce the z value for this gamma-ray
event. Low z values represent gamma-rays which have been scattered
and have lost a portion of their energy. These events are generally
not desirable for inclusion in an image because they carry false
position information. Therefore, the degree of rejection of each
event can be decided by software on the basis of the z value and a
spectrum of the number of gamma-rays versus the z value (energy)
can be recorded. This filtering process can be repeated for each
frame and all the frames can be added together to form the final
image. The operator can optionally go back to each original frame,
use a different z value threshold and reconstruct the final image
using different filter parameters. Variations in the sensitivity of
each pixel or superpixel can be mapped and included in the counter
for pixel by pixel corrections. The ability to discriminate
different radiation sources measured simultaneously or sequentially
includes defining filter parameters as selected energy threshold
values or ranges.
[0132] In this radionuclide imaging technique, the degree of
scatter rejection can be varied after the image acquisition in
order to decide on the optimal scatter rejection. This is not
possible with the conventional radionuclide imaging technology
employing a gamma camera or a rectilinear scanner. A gamma camera
or rectilinear scanner is generally incapable of detecting and
processing high intensity x-rays which are employed for high
quality x-ray radiography.
[0133] If an image intensifier is not used, the scintillator can be
in direct contact with the CCD. Alternatively, a fiber optic
reducer can be used between the CCD and the scintillator. Typical
reduction ratios vary from 1:1 to 6:1 although the present
embodiment is not limited to these ratios. Therefore, for a 20
mm.times.20 mm CCD, and a 6:1 fiber optic reducer, the area of
coverage will be about 120 mm. With a gated image intensifier or a
shutter, the CCD does not receive any signal during the readout
process. In a direct contact configuration, the use of frame
transfer CCD as shown in FIG. 10 is preferred.
[0134] In applications utilizing x-ray transmission measurements a
single frame is acquired for the recording of the x-rays emerging
from the irradiated body of tissue. The CCD is operating in the
integrating mode and each pixel or superpixel which accumulates a
charge which is proportional to the total number of x-rays in this
region without any energy discrimination. The resulting
radiographic image can be combined electronically with the
radionuclide image to form an accurate representation of both
physiologic and anatomic information.
[0135] In the case of thin specimens examined in vitro a light
source with wavelength ranging from the ultraviolet to near
infrared can be used for the transmission image in the integrating
mode. In this approach, the light shield in front of the
scintillator is removed and the detector is placed in an enclosure
to shield it from ambient light.
[0136] The present invention can thus combine radionuclide emission
imaging and x-ray transmission imaging (radiography) using the same
area detector with spectroscopic capability in the gamma-ray
imaging mode. This camera can be operated utilizing both, the
counting pulse-height analysis for gamma-ray imaging, and in the
integrating or counting modes for x-ray substantially transmission
imaging. This enables exact superposition of the two images for
accurate anatomic and physiologic imaging. Also, the operator can
change the energy threshold even after the radionuclide image has
been acquired. Thus, higher intrinsic spatial and energy resolution
are provided than found in the conventional approaches.
[0137] FIG. 15 is a schematic illustration of one preferred
embodiment of a dual-energy bone densitometry system 200 in
accordance with the invention. An x-ray tube 12 emits x-rays 14
which pass through the x-ray transparent patient table 254 and into
the patient (not shown). The x-rays 15 which pass through the
patient are directed through an x-ray transparent mirror 202 and
strike a first scintillator screen 204. The scintillator 204 reacts
to low-energy x-rays and generates a light pattern corresponding to
the low energy x-ray pattern. The light generated by the
scintillator 204 propagates back to the mirror 202 which reflects
the light to the lenses 206. The lenses 206 couple the image from
the scintillator 204 to an image intensifier 208 having
microchannel plates 210. Alternatively, the image intensifier 208
can be a proximity-type intensifier without the microchannel plates
210. The light from the image intensifier 208 is received and
detected by the detector 212, which can be a CCD array, a CID array
or an amorphous silicon sensor. The detector 212 senses the image
which corresponds to the low-energy x-rays and generates an
electronic representation of the image in the form of pixel
data.
[0138] High-energy x-rays pass through the scintillator 204 to an
optional x-ray filter 214. The filter 214 is preferably a copper
filter which blocks any remaining low-energy x-rays which pass
through the scintillator 204. An optional light block filter 216
can also be included between the scintillator 204 and the x-ray
filter 214 to block any stray optical radiation emanating from the
scintillator 204 from reaching a second detector 220.
[0139] The high-energy x-rays from the filter 214 strike a second
scintillator 218 which is reactive to the high-energy x-rays to
generate an optical image which corresponds to the pattern of
high-energy x-rays. The optical image is received by a second
detector 220, which can also be a CCD or CID array or an amorphous
silicon image sensor. The second detector 220 senses the optical
image and generates an electronic representation of the high-energy
x-ray pattern. An optional x-ray absorbing fiber optic plate 222
can also be included between the scintillator 218 and the detector
220 to absorb any remaining x-rays and thus prevent them from
interfering with the detector 220.
[0140] The system 200 of FIG. 15 can be used in either a scanning
mode or a stationary mode. In the scanning mode, the x-ray tube
source 12 as well as the detection system are moved continuously or
in a stepping motion along the region being examined. While the
system scans the region, a series of images are obtained having
short exposure acquisition times. In the stationary mode, a single
exposure is made of the entire region being examined. Time delay
integration (TDI) is used in which the CCD stores the total charge
for each pixel during a selected x-ray exposure interval. At the
end of the x-ray exposure, the discrete representation in each
pixel is readout by a CCD controller. Once the data is thus
obtained, the comparative processing techniques of dual photon
absorptiometry can be used to determine quantitative density
measurements of the calcified material such as bone within the body
regions exposed by the x-rays.
[0141] In the system 200 of FIG. 15, the image intensifier 208 can
be omitted. In that configuration, to ensure that image data for
the low-energy x-rays can be accurately collected, the detector 212
can be cooled to increase signal-to-noise ratio.
[0142] FIG. 16 is a schematic diagram of another embodiment of a
dual-energy bone densitometry measuring system 300 in accordance
with the invention. An x-ray tube 12 outputs x-rays 14 through
x-ray transparent patient table 254 and into the patient. X-rays 15
directed through the patient strike a first scintillator 302 which
is reactive to low-energy x-rays to generate an optical image of
the low-energy x-ray pattern out of the patient. The optical image
is carried by a coherent fiber optic conduit 304 to a CCD detector
306 which detects the optical image and generates the electronic
representation of the low-energy x-ray pattern. The fiber optic
conduit 304 is preferably made of plastic optical fibers to
facilitate collection of the low-energy image. However, if the
distance labeled "x" is selected to be small enough, glass fibers
can be used instead. The space labeled 310 is filled with a film
material being the same material as that of which the fibers are
made.
[0143] The high-energy x-rays pass through the scintillator 302,
the fiber optic conduit 304 and the film material 310 and strike a
second x-ray phosphor scintillator 312. The second scintillator 312
is reactive to high-energy x-rays and therefore generates an
optical image which corresponds to the high-energy x-ray pattern.
The optical image generated by the scintillator 312 is detected by
a second CCD array 314 which generates the electronic
representation of the high-energy x-ray pattern. An optional copper
or aluminum filter 316 can be inserted in front of the second
scintillator 312 to absorb any remaining low-energy x-rays. Also,
an x-ray absorbing fiber optic plate 308 can be inserted between
the scintillator 312 and the CCD 314 to prevent x-rays from
impinging on the CCD 314.
[0144] FIG. 17 is a schematic diagram of another embodiment of a
dual-energy bone densitometer measuring apparatus 400 in accordance
with the invention. The system 400 of FIG. 17 is the same as the
system 300 of FIG. 16 except that the coherent fiber optic conduit
304 in FIG. 16 is replaced with a different conduit 404 in the
system 400 of FIG. 17. In the conduit 404 of FIG. 17, the fibers
are bent at approximate right angles with small radii of curvature.
As in the embodiment of FIG. 16, the fibers are either plastic or
glass. Because of the different fiber bending in which the
collected radiation is redirected from a first optical path onto a
second optical path, the need for the film material 310 shown in
FIG. 16 is eliminated.
[0145] FIG. 18 is a schematic diagram of another embodiment of a
bone densitometer measuring apparatus 500 in accordance with the
invention. In this embodiment, scintillator plates 505 and 507 are
used to convert the x-ray energy into optical energy. Once again,
the x-ray tube 12 directs x-rays 14 through the patient table 254
and the patient. The x-rays 15 emanating from the patient first
strike an anti-scatter grid 502 which prevents scattered x-rays
from reaching the detectors. The x-rays then strike a first
amorphous silicon image sensor 504 which detects low-energy x-rays
and generates the data which indicates the low-energy x-ray
pattern. The low energy sensor 504 can be thinner than the
high-energy sensor 508 to reduce the filtering requirements of the
system. Also scintillator 505 can be thinner than scintillator 507
to improve collection efficiency of the system. High-energy x-rays
pass through the first sensor 504 and then through a copper,
tungsten, gadolinium or aluminum x-ray filter 506 which filters out
low-energy x-rays. The high-energy x-rays then strike the second
amorphous silicon image sensor 508 which generates the data for the
high-energy x-ray pattern. The low-energy x-ray pattern data and
the high-energy x-ray pattern data are read out of the amorphous
silicon image sensors 504 and 508, respectively, by a detector
controller 510.
[0146] FIG. 19 is a schematic diagram of an alternative detection
structure 550 which can be used with the dual-energy bone
densitometry measuring apparatus 500 of FIG. 18. The lower layer of
the structure 550 is an anti-scatter grid 552 used to prevent
scattered x-rays from reaching the detection structure 550. The
next layer is a low energy x-ray scintillator layer 554 which
generates an optical image of the low-energy x-ray pattern. An
amorphous silicon image sensor 556 detects the optical image from
the scintillator 554 to generate the data for the low-energy x-ray
pattern. A substrate layer 558 is formed over the amorphous silicon
image sensor layer 556. The substrate layer 558 includes a thinned
central region 560. The thinned substrate 558 provides for
increased transmission to the second scintillator layer 562. The
second scintillator 562 is reactive to high-energy x-rays to
generate an optical image of high-energy x-ray pattern. The optical
image is detected by a second amorphous silicon image sensor 564.
The structure 550 is covered by a protective substrate 566,
preferably made of glass. A thin layer of lead can be formed on top
of the glass to prevent propagation of x-rays beyond the structure
550. Preferred scintillators include C.sub.s(+1), C.sub.dW0.sub.4,
or gadolinium oxysulfide.
[0147] The amorphous silicon array sensors and the associated
control and processing systems can utilize the binning and other
processing capabilities described elsewhere in the present
application. Additionally, a plurality of such sensors can be
combined to form a single or dual array. The array can be linear,
rectangular or square depending upon the particular application.
The systems can be used in conjunction with a C-arm assembly where
the C-arm 580 rigidly aligns the source 586 and detector assembly
582 as shown in FIG. 20. The C-arm 580 can also be used to rotate
the source and detector about the patient on table 584 as indicated
at 588 to provide multidirectional viewing of the entire human
skeletal structure including the hip and femur. Thus lateral spine
imaging and quantitative analysis can be conducted using the
present system. The detector assembly 582 includes a CCD sensor as
described herein in conjunction with a straight or angled fiber
optic coupler and scintillator. The detector assembly 582 can be
scanned or stepped along axis 590 and axis 592 that is parallel to
the spine of the patient in order to provide a sequence of images
for both quantitative and qualitative analysis.
[0148] The detector assembly 582 can include various configurations
described elsewhere herein, including the examples illustrated in
FIG. 21A, 21B, and 21C. In FIG. 21A a straight fiber optic coupler
602 optically couples the scintillator 604 to the CCD (or CID or
amorphous) sensor array 600. An optional cooler(s) 606 can be used
in these examples. In FIG. 21B a fiber optic reducer 608 couples
the scintillator 610 to the sensor array 600. A proximity type
x-ray image intensifier and scintillator can replace scintillators
604 and 610. In FIG. 21C a dual sensor system includes sensors 600
and 612, fiber optic coupler 602, scintillators 618, 620, mirror
616, and lens 614. This system functions in a manner similar to
that described in connection with FIG. 15.
[0149] FIGS. 22A and 22B illustrate a preferred method of imaging
in which the entire imaging field is composed of a series of
slightly overlapping individual images 620 that are acquired by a
continuous scan or stepped imaging sequence along the rectilinear
path 622. Dual energy tissue or bone density measurements can be
accomplished by collecting data at two energies at each subfield
620. The x-ray source can be switched or filtered as described
previously to generate discrete energy peaks.
[0150] FIG. 23 illustrates a fan-beam system in which the x-ray
source 586 generates a fan shaped beam 640 that is detected by a
detector system 700. System 700 can include a scintillator, fiber
optic plate or reducer for each of a plurality of sensors 630 which
are aligned in a linear array to collect fan beam 640. Detector
system 700 can use a lead slit collimator 702 and can use CCDs,
CIDs or a number of amorphous silicon sensors in configurations
illustrated, for example, in FIGS. 21A-21C.
[0151] FIG. 24 illustrates another preferred embodiment 650 in
which a patient 654 is positioned on table 652. X-ray tube 656
directs fan-beam 660 through a scanning slit collimator 658, the
patient 654 and a second scanning slit collimator 664. The
radiation 660 then passes through mirror 62 striking the
scintillator 676. The scintillator emits light that is reflected by
mirror 662 towards the sensor 672 as illustrated at 674. Optional
lead glass element 666 can be placed at any position between the
mirror 662 and the sensor 672. A lens 668 and cooler 670 can also
be employed, if necessary. Lead foil 678 can be used to line the
enclosure 710 to reduce interactions between the scattered x-rays
and the sensor 672. The system can alternatively use a proximity
type image intensifiers in the x-ray path before the mirror.
[0152] FIG. 25 is a schematic diagram of a detection system 800
which can be used with the systems described above for dual-energy
bone densitometry measurements as well as tissue and lesion
imaging. The system 800 can include an enclosure 802 having an
aperture 804 through which radiation such as x-ray beams 806 enter
the system 800. In one embodiment, the x-ray beams 806 pass through
an x-ray transparent mirror 808 and strike a first scintillating
plate 809. The first scintillator 809 is reactive to low-energy
x-rays and produces an optical image corresponding to the
low-energy x-ray pattern. The optical image is projected back onto
the mirror 808 which reflects the image to lens 812. The lens 812
focuses the light onto a first surface of a CCD array detector 814.
The detector 814 can include a proximity-type image intensifier to
enhance image detection capabilities. An annular cooler can also be
placed around the CCD detector 814 to cool the CCD and therefore
improve signal-to-noise ratio.
[0153] High-energy x-rays pass through the top scintillator 809 and
strike the lower second scintillator 810 which is reactive to the
high-energy x-rays to produce an optical image which corresponds to
the high-energy x-ray pattern. The optical image is reflected by a
second mirror 816 to a third mirror 818 which directs the light
through a second focusing lens 820. The lens 820 focuses the light
onto the back surface of the CCD detector 814. The back surface can
also include a proximity-type image intensifier. In addition, the
annular cooler, if present, cools the front and back surfaces of
the CCD detector 814.
[0154] Thus, the system 800 of FIG. 25 produces increased
sensitivity by sensing on opposite sides of a single thinned CCD
detector 814. Spatial correlation between the two images which are
fused to form a single image is greatly improved over the
previously described embodiments with separate detection surfaces
since the relative locations of the image detection surfaces can be
more precisely controlled. High and low energies can be detected on
both sides.
[0155] The detection system 800 of FIG. 25 can include plural
two-sided CCD detectors to provide the system with a wide
field-of-view along with combined electronics and cooling. FIG. 25
illustrates a second detector 834 and associated lenses 830 and
832. It will be understood that more detectors and lenses can be
added as needed.
[0156] The dual-energy configuration of the system 800 described
above facilitates bone densitometry measurements as previously
described. However, the system 800 can also be used for detecting
and imaging lesions in patient tissue as described above. In that
embodiment, the x-ray beams 806 are replaced with other types of
radiation such as in the visible or infrared ranges. The
scintillators 809 and 810 and the mirrors 808, 816 and 818 can be
used to form images of the tissue to be formed at two different
wavelengths on opposite surfaces of the detectors 814 and 834.
[0157] A particular application of the methods and systems
described herein for detecting and imaging of soft tissue lesions
includes digital mammography using CCDs or similar type
silicon-based detectors such as amorphous silicon type detectors
described above. These systems are used to detect lesions in the
tissue including calcified material within the soft tissue, that
can indicate the need for more careful diagnostic procedures and/or
treatment of the patient. Slot-scanning approaches using time-delay
integration, where the CCD records continuously during a scan as
described herein, can be used for digital mammography. However, the
continuous recording approach results in certain problems,
particularly with artifacts due to the shear distortion of the
fiber-optic plates which can be used with such an embodiment. While
slot-scan approaches using the continuous record mode can be used,
the quality of the images is less than ideal due to the distortion
effects.
[0158] Other methods for scanning the breast include the dividing
of the image area into four quadrants or even a greater number of
segments. Every time it is necessary to take multiple exposures of
the breast, the associated problems including increased exposure
level and collection time can limit the variety of applications for
which the system can be used. It is desirable, therefore, if one
needs to acquire the image in a step-wise fashion, that there be no
more than two, or at most, three acquisition steps. If one uses a
greater number of acquisition steps, the breast has to remain
compressed for too long, thus causing extreme discomfort to the
patient. Moreover, the x-ray tube power requirements increase
significantly. The preferred method for digital mammography
applications involving sequential multiple imaging is thus limited
to a two image acquisition process. This procedure involves
directing x-rays from the source through the tissue to the
stationary detector system for about 0.2-5.0 seconds and preferably
in the range of 0.5-1 second, the detector system can then be moved
to a second position while the first image is read out, then a
second exposure is obtained and read out. As many as 2-5 million
pixels can be read out in a time interval that is less than the
exposure interval.
[0159] A problem associated with the two-dimensional array
approach, is its complexity and cost. Although the tiling of
4.times.3 CCDs, for example, to form an array can be used for
digital mammography, it is likely to be too expensive for many
common applications. This results from the cost of the CCDs
themselves, and with the problems associated with making a seamless
joint to three or four sides of the CCDs.
[0160] Referring to FIGS. 26-32, the detector module 900 can
consist of from 3 to 5 CCDs in a first linear array 902 and another
set of CCDs can be positioned approximately 6 cm apart in a second
linear array 904. This embodiment utilizes four CCD elements in
each array. Each element can include a scintillator 906 and a
tapered fiber optic plate 908. In the embodiment where the CCDs are
replaced by amorphous silicon sensors, a single strip silicon
sensor can be substituted for each linear array in these
embodiments.
[0161] The first set of CCDs can be placed as closely as possible
to the chest wall of the patient. The x-ray beam is collimated by
using a double slot to provide two fan beams, each fan beam being
directed onto a linear array, thus only two areas are irradiated
which correspond exactly to each CCD group. After one x-ray
exposure and acquisition, the x-ray collimator is translated in
synchrony with both CCD banks which are also translated to the next
position. Another exposure is taken and the signal is read out. A
small amount of overlapping of the fields, about 1-3 mm can be
desirable. With the use of a micro-stepping translation stage the
successive fields can be aligned to within a few microns with or
without overlap. The images can then be joined and will be
substantially seamless with less than a 5-10 micron difference
between the region of the body and the joined images of the
region.
[0162] The sensing surface does not have to be on a plane. As shown
in FIG. 27, the CCDs 910 can be arranged on a curving or non-planar
surface. This is an extremely important embodiment because it
provides for the use of straight (non-tapering) fiber-optic plates
912 which dramatically reduces the cost and contributes to better
image quality. Please note that the CCDs can be cooled or
non-cooled and can be operated in the pixel binned or non-binned
mode. Additionally, an anti-scatter grid can be used between the
breast and the detectors. Each element 910 in the array is
generally equidistant from the x-ray source in order to reduce
distortion across the entire field of view of the array. This arced
linear array can be used for many different applications as
described elsewhere herein.
[0163] This approach is preferable as current manufacturers can
readily make CCDs which are buttable on two sides. It remains
difficult and expensive to make CCDs buttable on three or four
sides. In the illustrated embodiment there are only six joints
required between the CCDs, unlike a large area cassette which has
many more. A typical CCD for this application can have an area of
6.times.6 cm but for economy reasons, one can use a larger number
of CCDs, such as 3.times.3 cm elements. For example, if one uses a
3.times.3 cm device, each CCD linear array 902, 904 incorporates
eight CCDs for a total of 16 CCDs. This can also be used to provide
a larger area of coverage comparable to a standard large film
cassette. FIG. 28 illustrates a four line separated array 916
covering a 24.times.30 cm planar area with 10 of the 3.times.3 cm
devices in each line 918. In FIG. 29, a partial cross sectional
view of one of the lines 918 in FIG. 28, illustrates a preferred
embodiment in which each CCD 914 is butted against one or two
adjoining CCDs in each line with each line coupled to a
scintillator 915 and a fiber optic plate 917. The two-step
acquisition is preferable relative to the narrow slot-scanning
approaches which typically use a slot width of about 1.5 cm, and
the larger area imaging approaches which are effective but can be
extremely costly.
[0164] As illustrated in FIG. 30, an x-ray source 922 and a double
or multiple slot collimator 924 can be used to generate and align
the x-rays 928 with the translating CCD modules. An actuator or
motorized system 920 is used to translate both CCD arrays 902, 904
without altering the distance between the rigidly aligned CCD
arrays. The system 920 can be connected to a controller or personal
computer as described previously so that the user can control array
position along the direction of translation 926, which in this
embodiment, is towards or away from the chest wall.
[0165] As shown in FIGS. 31A and 31B, the arrays 902, 904 are
positioned to image two parallel regions 930, 932. The detectors
902, 904 are then translated from the first position to a second
position to image and analyze two further parallel regions 934, 936
to provide full image of compressed breast 925. The relative
spacing between the two linear arrays can also be controlled to
increase or decrease overlap. A preferred embodiment, however,
retains the two spaced arrays in a rigid position relative to each
other. This particular embodiment moves the detectors towards or
away from the chest wall of the patient.
[0166] Shown in FIG. 32 is an embodiment 940 in which the array 916
of FIG. 28 in which the direction of scan 942 is along the chest
wall. The collimator 944 is also moved along the same axis as the
array to direct x-rays 928 onto the CCDs 948 and not onto the
spaces 946.
[0167] While the invention has been particularly shown and
described with reference to a preferred embodiment thereof, it will
be understood by those skilled in the art that various changes in
form and details can be made therein without departing from the
spirit and scope of the invention as defined by the appended
claims.
* * * * *