U.S. patent application number 09/119144 was filed with the patent office on 2002-03-28 for method and apparatus for gamma ray detection.
Invention is credited to TUMER, TUMAY O..
Application Number | 20020036270 09/119144 |
Document ID | / |
Family ID | 46276254 |
Filed Date | 2002-03-28 |
United States Patent
Application |
20020036270 |
Kind Code |
A1 |
TUMER, TUMAY O. |
March 28, 2002 |
METHOD AND APPARATUS FOR GAMMA RAY DETECTION
Abstract
A high sensitivity, three-dimensional gamma ray detection and
imaging system is provided. The system uses the Compton double
scatter technique with recoil electron tracking. The system
preferably includes two detector subassemblies; a silicon
microstrip hodoscope and a calorimeter. In this system the incoming
photon Compton scatters in the hodoscope. The second scatter layer
is the calorimeter where the scattered gamma ray is totally
absorbed. The recoil electron in the hodoscope is tracked through
several detector planes until it stops. The x and y position
signals from the first two planes of the electron track determine
the direction of the recoil electron while the energy loss from all
planes determines the energy of the recoil electron.
Inventors: |
TUMER, TUMAY O.; (RIVERSIDE,
CA) |
Correspondence
Address: |
RONALD R. SNIDER
P.O. BOX 27613
WASHINGTON
DC
200387613
|
Family ID: |
46276254 |
Appl. No.: |
09/119144 |
Filed: |
July 20, 1998 |
Related U.S. Patent Documents
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
|
|
09119144 |
Jul 20, 1998 |
|
|
|
08784176 |
Jan 15, 1997 |
|
|
|
5821541 |
|
|
|
|
60011135 |
Feb 2, 1996 |
|
|
|
Current U.S.
Class: |
250/370.09 |
Current CPC
Class: |
A61B 6/4258 20130101;
G01T 1/1647 20130101; G01T 1/2928 20130101; G01T 1/006
20130101 |
Class at
Publication: |
250/370.09 |
International
Class: |
G01T 001/24 |
Claims
What is claimed is:
1. A method of imaging a portion of a human body, the method
comprising the steps of: providing a radiopharmaceutical to said
portion of said human body, said radiopharmaceutical producing
gamma ray photons; positioning a detection system proximate to said
portion of said human body, wherein said detection system is
comprised of a hodoscope and a calorimeter, wherein said hodoscope
is comprised of a plurality of detector planes; determining a
direction and an energy for a portion of said gamma ray photons
entering said detection system from a plurality of hodoscope output
signals and from a plurality of calorimeter output signals;
processing said direction and energy data for said portion of gamma
ray photons; and displaying an image of said portion of said human
body, wherein said image is based on said processed direction and
energy data.
2. The method of claim 1, wherein said portion of said human body
is an organ.
3. The method of claim 1, wherein said portion of said human body
is a breast.
4. The method of claim 1, further comprising the step of
positioning a shielding member around an entrance of said detection
system, wherein said shielding member limits background gamma ray
photons.
5. The method of claim 1, further comprising the steps of:
positioning a second detection system proximate to said portion of
said human body, wherein said second detection system is comprised
of a second hodoscope and a second calorimeter, wherein said second
hodoscope is comprised of a second plurality of detector planes;
determining a second direction and a second energy for said portion
of gamma ray photons entering said second detection system from a
plurality of second hodoscope output signals and from a plurality
of second calorimeter output signals; processing said second
direction and second energy data for said portion of gamma ray
photons; and displaying a three-dimensional image of said portion
of said human body, wherein said image is based on said processed
first and second direction data and said first and second energy
data.
6. The method of claim 1, further comprising the steps of:
repositioning said detection system proximate to said portion of
said human body; determining a second direction and a second energy
for said portion of gamma ray photons entering said detection
system; processing said second direction and second energy data for
said portion of gamma ray photons; and displaying a
three-dimensional image of said portion of said human body, wherein
said image is based on said processed first and second direction
data and said first and second energy data.
7. The method of claim 1, wherein said processing step comprises
applying a data analysis technique selected from the group
consisting of a Radon transform and back projection technique, a
Maximum Likelihood and Maximum Entropy technique, a Direct Linear
Algebraic Deconvolution technique, and a Constrained Linear
Algebraic Deconvolution technique.
8. A medical imaging system for imaging a portion of a living
organism, said portion treated with a radiopharmaceutical, said
radiopharmaceutical emitting gamma ray photons, comprising: a
hodoscope comprised of a plurality of silicon detection planes,
wherein an entrance aperture of said hodoscope is external to said
living organism and proximate to said portion of said living
organism, wherein emitted gamma ray photons pass into said
hodoscope and are scattered within said hodoscope; a multi-channel
readout system coupled to said plurality of silicon detection
planes; a processor coupled to said multi-channel readout system;
and a monitor coupled to said processor, said monitor displaying an
image of said portion of said living organism.
9. The medical imaging system of claim 8, wherein a portion of said
emitted gamma ray photons undergo at least one Compton scatter
within said plurality of silicon detection planes to yield a track
direction corresponding to each of said portion of emitted gamma
ray photons, wherein a total energy corresponding to each of said
portion of said emitted gamma ray photons is absorbed within said
plurality of silicon detection planes, and wherein said track
direction and said total energy corresponding to each of said
portion of emitted gamma ray photons is combined by said processor
to generate said image.
10. The medical imaging system of claim 8, wherein a recoil
electron is formed by a portion of said emitted gamma ray photons
undergoing Compton scatter within said plurality of silicon
detection planes, said recoil electron passing through a portion of
said plurality of silicon detection planes, wherein a position of
said recoil electron is recorded for each of said portion of said
plurality of silicon detection planes.
11. The medical imaging system of claim 8, further comprising a
calorimeter enclosing a portion of said hodoscope, wherein said
calorimeter is coupled to said multi-channel readout system.
12. The medical imaging system of claim 11, wherein said scattered
gamma ray photons passing through said hodoscope form recoil
electrons during passage through said plurality of silicon
detection planes, wherein said hodoscope determines a track
direction by a first scatter vertex and an energy associated with
said recoil electrons, wherein said scattered gamma ray photons are
totally absorbed within said calorimeter, and wherein an energy of
said absorbed gamma ray photons is determined by said
calorimeter.
13. The medical imaging system of claim 8, wherein said
radiopharmaceutical is selected from the group consisting of
thallium-201, technetium-99m, iodine-123, iodine-131, and
fluorine-18.
14. The medical imaging system of claim 11, further comprising a
shielding member proximate to said hodoscope entrance aperture
positioned to limit background gamma ray photons.
15. The medical imaging system of claim 8, wherein said plurality
of silicon detection planes is selected from the group consisting
of silicon microstrip detectors, silicon strip detectors, silicon
pad detectors, silicon pixel detectors, double sided silicon
microstrip detectors, and double sided silicon strip detectors.
16. The medical imaging system of claim 8, wherein said plurality
of silicon detection planes have a predetermined orientation with
respect to said entrance aperture, said predetermined orientation
selected from the group consisting of parallel, perpendicular, or
an angle.
17. The medical imaging system of claim 8, wherein each of said
plurality of silicon detection planes is comprised of at least two
bridged silicon detectors, wherein said silicon detectors are
selected from the group consisting of strip and microstrip
detectors.
18. The medical imaging system of claim 8, wherein each of said
plurality of silicon detection planes has a corresponding thickness
within a range of about 100 micrometers to about 5 millimeters.
19. The medical imaging system of claim 8, wherein said plurality
of silicon detection planes are separated by a distance of between
about 0.2 and 2 centimeters.
20. The medical imaging system of claim 8, wherein said plurality
of silicon detection planes is between 10 and 25 silicon detection
planes.
21. The medical imaging system of claim 11, wherein said
calorimeter is selected from the group of calorimeter detector
materials consisting of HPGe, BGO, CdWO.sub.4, CsF, NaI(Tl),
CsI(Na), CsI(Tl), CdTe, CdZnTe, HgI.sub.2, GaAs, and PbI.sub.2.
22. The medical imaging system of claim 11, wherein said
calorimeter is comprised of CdZnTe detectors selected from the
group consisting of CdZnTe strip detectors, CdZnTe pad detectors,
and CdZnTe pixel detectors.
23. The medical imaging system of claim 11, wherein said
calorimeter is comprised of CsI(Tl) crystals coupled to PIN
photodiodes.
24. The medical imaging system of claim 11, wherein said
calorimeter is comprised of multiple calorimetry layers.
25. The medical imaging system of claim 11, wherein said portion of
said hodoscope enclosed by said calorimeter is a side portion.
26. The medical imaging system of claim 11, wherein said portion of
said hodoscope enclosed by said calorimeter is a back portion,
wherein said back portion is opposed to said hodoscope entrance
aperture.
27. The medical imaging system of claim 8, said processor applying
a data analysis technique selected from the group consisting of a
Radon transform and back projection technique, a Maximum Likelihood
and Maximum Entropy technique, a Direct Linear Algebraic
Deconvolution technique, and a Constrained Linear Algebraic
Deconvolution technique.
28. The medical imaging system of claim 8, wherein said emitted
gamma ray photons have an energy within a range of about 40 keV to
about 2,000 keV.
29. The medical imaging system of claim 11, further comprising a
slot collimator proximate to said hodoscope entrance aperture.
30. The medical imaging system of claim 8, wherein said
multi-channel readout system is comprised of ASIC chips, wherein
said plurality of silicon detection planes is comprised of silicon
strip detectors, and wherein said strips are fanned in to match a
chip bonding pitch corresponding to said ASIC chips.
Description
CROSS-REFERENCES TO RELATED APPLICATIONS
[0001] This application is a continuation-in-part of U.S.
application Ser. No. 08/784,176, filed Jan. 15, 1997 which is a
continuation of U.S. Provisional Application Serial No. 60/011,135,
filed Feb. 2, 1996 (now abandoned).
FIELD OF THE INVENTION
[0002] The present invention relates generally to detection
systems, and more particularly, to a method and apparatus for
imaging gamma rays.
BACKGROUND OF THE INVENTION
[0003] The various organs and tissues of the human body fall prey
to a myriad of different afflictions. For example, each year in the
United States alone, approximately 180,000 women are diagnosed with
breast cancer and 46,000 women die of this disease. In all, 10 to
11 percent of all women can expect to be affected by breast cancer
at some time during their lives. The causes of most breast cancers
are not yet understood. Screening and early diagnosis are currently
the most effective ways to reduce mortality from this disease.
[0004] Currently mammography is the most effective means of
detecting non-palpable breast cancer. However, mammography cannot
determine whether a lesion is benign or cancerous, typically one or
more biopsies must be performed per lesion. Unfortunately the
biopsy operation itself is a very traumatic and costly operation
that often results in some degree of disfigurement. Therefore it is
important to improve the specificity of mammography thereby
reducing errors, patient trauma, and disfiguration from unnecessary
biopsies. It is also important to reduce health care costs by
decreasing the number of unnecessary biopsies. For example, to
detect 100,000 non-palpable cancers, approximately 500,000 biopsies
must be performed at a cost of about $5,000 per biopsy, yielding a
total cost of approximately 2.5 billion dollars. Therefore a
reduction of 50 percent would save about 1.25 billion dollars per
year.
[0005] Palpable mass abnormalities of the breast are often
difficult to evaluate mammographically. This is especially true for
patients with dense or dysplastic breasts (approximately 35 percent
of women over 50 and 70 percent of women under 50) or those
patients that exhibit signs of a fibrocystic change, for example
due to radiation therapy. For example, invasive lobular carcinoma
in dense breasts can attain a size of several centimeters and yet
still show no mammographically detectable signs. Furthermore, about
50 percent of all preinvasive cancers do not show mammographically
significant calcifications, thus decreasing the chances of
detecting the malignant tumors.
[0006] Lastly, due to the interpretational limitations of
mammography, many high risk patients (i.e., patients with a family
history of breast cancer, patients with prior histologic evidence
of cellular atypia, patients with a prior history of breast cancer
who have undergone lumpectomy and radiation therapy) may be forced
to rely on random tissue biopsies performed on suspicious areas.
Unfortunately this technique typically results in a high
nonmalignant-to-malignant biopsy ratio.
[0007] A relatively new scientific tool that has allowed scientists
and physicians to address problems in physiology and biochemistry
in the human body with low risk is emission computed tomography
(ECT). ECT systems are mainly used for the detection and imaging of
the radiation produced by radiotracers and radiopharmaceuticals.
For example, by administering biologically active
radiopharmaceuticals into a patient it is possible to image organ
functions in real time.
[0008] The two major instruments presently used for ECT are Single
Photon Emission Computed Tomography (SPECT) and Positron Emission
Tomography (PET). These instruments have been used to study a
variety of different organs and conditions including cerebral
glucose consumption, protein synthesis evaluation, cerebral blood
flow and receptor distribution imaging, oxygen utilization, stroke,
heart, lung, epilepsy, breast cancer, dementia, oncology,
pharmacokinetics, psychiatric disorders, and radio labelled
antibody and cardiac studies. Since the SPECT and PET instruments
use different types of radiotracers, the metabolic activities
imaged are mostly different leading these two instruments to
complement rather than compete with each other. The SPECT detectors
have proven especially useful for heart and brain imaging.
[0009] SPECT dates back from the early 1960s, when the first
transverse section tomographs were presented by Kuhl and Edwards
(1963) using a rectilinear scanner and analog back-projection
methods. With the availability of computer systems and the impetus
of computer-assisted tomography using transmitted x-rays, nuclear
medicine instruments were modified, and a number of mathematical
approaches to tomographic reconstruction were developed in the
early 1970s. Rotating Anger cameras and advances in computers
opened the way to three-dimensional SPECT systems. Recently
interest in SPECT increased as mathematical reconstruction
techniques improved. They allowed for attenuation compensation,
scattered radiation correction and the availability of new
radiopharmaceuticals with higher uptake in the brain or other
organs. The major limiting factors for the SPECT systems presently
are the sensitivities (.apprxeq.10 Cts s.sup.-1 .mu.Ci.sup.-1 point
and .apprxeq.1,000 Cts s.sup.-1 cm.sup.-1 volume), resolution (7 to
12 mm FWHM), size, and cost.
[0010] Present SPECT systems mainly use the rotating Anger camera.
Many different variations of the Anger camera and other smaller
size rotating single or dual instruments have been designed and
used. Most of the commercial instruments use NaI(Tl), CsI(Tl), CsF,
BaF.sub.2, BGO and other related crystal detectors. The majority of
the commercial instruments use the Anger cameras made of NaI(Tl)
crystals. All commercial SPECT instruments use collimators for
determination of the direction of the incident gamma rays. The main
types are parallel and converging collimators. The converging fan
or cone beam collimators produce higher sensitivity but increase
the complexity of the data analysis. Pinhole and slit collimators
are also used. The collimators for high resolution systems
eliminate about 99.9 percent of the incident gamma rays. A typical
collimator hole has an area of about 1 square millimeter and a
length of 1.9 centimeters. Increasing collimator resolution
decreases sensitivity and vice versa. Collimators made of high
atomic number materials such as lead which also produce
considerable amounts of scattered gamma rays on the inside surface
of the collimator, thereby increasing the scattered photon
background.
[0011] Anger cameras are normally rotated on a gantry around the
patient for about 20 minutes to acquire sufficient data for a
reasonable image. The spatial resolutions are limited to about 7 to
12 millimeters although spatial resolutions are expected to reach 6
millimeters in the near future. The best energy resolution at gamma
ray energies is about 10 percent, limiting the ability of Anger
cameras to discriminate scattered photon background. Commercially
available SPECT systems include ADAC ARC, GE Starcam, Elscint APEX,
Trionix Triad, Digital Scintigraphics ASPECT and University of
Michigan SPRINT II.
[0012] From the foregoing it is apparent that an improved gamma ray
imaging system is desired.
SUMMARY OF THE INVENTION
[0013] The present invention provides a high sensitivity, high
spatial resolution, and electronically collimated single photon
emission computed tomography (SPECT) system. Its primary
sensitivity is in the range of 81 keV to 511 keV although it can be
used to detect higher energies of up to a few MeV by increasing the
detector thickness for both the hodoscope and the calorimeter. Both
the direction and energy of the incident gamma ray photons is
measured with high resolution. The method of determination of the
photon direction eliminates the need for a mechanical collimator
and the energy measurement discriminates against the scattered
photon background.
[0014] The disclosed system is constructed from position sensitive,
double sided silicon strips with a strip pitch of approximately 1
millimeter or silicon microstrips with a strip pitch much less than
a millimeter. Preferably the system uses the silicon strip
detectors. These detectors, varying in thickness from 150
micrometers to 2 millimeters, can produce the x and y coordinates
of a photon interaction in a single wafer.
[0015] One embodiment of the system uses multiple planes of double
sided silicon strip detectors with about 1 millimeter pixel size
and a thickness of 100 micrometers to 5 millimeters. The planes are
separated by a distance of between 0.2 and 2 centimeters, depending
on the pixel size and the required angular resolution. The smallest
possible separation is always preferred to keep the depth of the
detector small without sacrificing spatial resolution. The incident
gamma ray Compton scatters in one of the detector planes, the
dominant process for photons with at least 50 keV energies in
silicon strip detectors. The energy of the scattered electron in
this detector plane is measured. The scattered gamma ray with
reduced energy can be absorbed in the calorimeter or in an another
detector plane through the photoelectric effect or undergo multiple
Compton scatters followed by a photoelectric effect. The energies
of these subsequent interactions are also measured. If the
scattered gamma ray photon is completely absorbed, the sum of the
two energies gives the energy of the incident photon and the
individual energies and direction of the scattered photon give the
scatter angle of the incident gamma ray. Thus the gamma rays
emitted from a radionuclide can be imaged without need for a
collimator.
[0016] The scattered gamma ray photons can make a second Compton
scatter and then escape without further interaction. Also the
photons already scattered inside the patient will deposit lower
total energy. These events will produce a tail at lower energies in
the energy spectrum. Such events can be discriminated effectively
because the total energy detected is smaller than the known
incident gamma ray energy. However, a high sensitivity mode may be
applied with reduced angular resolution by adding the missing
energy to the energy measured at the second scatter. This will
dramatically increase the sensitivity but reduce angular resolution
somewhat and will not allow the discrimination of the scattered
photon background.
[0017] A calorimeter surrounding the silicon strip detector
hodoscope absorbs the Compton scattered photons. The calorimeter
can be fabricated from a plane of silicon a few millimeters thick,
CdZnTe strip and/or detectors, or CsI(Tl) crystals viewed by a
photodiode. The calorimeter can also be used as a second scatterer
and/or a missing energy detector.
[0018] The double Compton scatter measurement determines the
direction of the incident gamma ray to a cone with a half angle
equal to the scatter angle. This type of measurement is new in
nuclear medicine and requires special data analysis software. The
data analysis can be carried out by cone interaction, Maximum
Likelihood or Maximum Entropy techniques. These are iterative
techniques and require long computation times. A new direct data
analysis and imaging technique, Direct Linear Algebraic
Deconvolution (DLAD) method, can also be applied for real time
imaging.
[0019] In use, the present system utilizes the higher uptake of
certain radiopharmaceuticals by the organ or tissue of interest,
thereby allowing the selected organ/tissue to be imaged. For
example, malignant tissues preferentially absorb Tc-99m SestaMIBI
and Tl-201 chloride as compared to benign masses (except for some
highly cellular adenomas). Therefore, these radiopharmaceuticals
can be used to help diagnose and differentiate tumors from benign
growths, for example in a scintimammography system for breast
cancer detection and diagnosis. Possible mechanisms for uptake of
Tl-201 chloride into tumor cells include the action of the ATPase
sodium-potassium transport system in the cell membrane which
creates an intracellular concentration of potassium greater than
the concentration in the extracellular space. Thallium may be
significantly influenced by this system in tumors. In addition, a
co-transport system has been identified which also is felt to be
important in uptake of thallium by tumor cells. The mechanism of
Tc-99m SestaMIBI accumulation in tumors is not yet.
[0020] A further understanding of the nature and advantages of the
present invention may be realized by reference to the remaining
portions of the specification and the drawings.
BRIEF DESCRIPTION OF THE DRAWINGS
[0021] FIG. 1 is an illustration of the gamma ray linear
attenuation coefficients for a silicon detector;
[0022] FIG. 2 is an illustration of the Compton scatter technique
for detecting gamma rays;
[0023] FIG. 3 is an illustration of a typical double sided silicon
microstrip or strip detector;
[0024] FIG. 4 is the schematic cross section of the detector
illustrated in FIG. 3;
[0025] FIG. 5 is an illustration of a side view of an embodiment of
the invention;
[0026] FIG. 6 is an illustration of a top view of the system
illustrated in FIG. 5;
[0027] FIG. 7 is an illustration of the FEE readout chips and the
silicon strip detector planes mounted on a printed circuit
board;
[0028] FIG. 8 is a graph illustrating the energy spectrum of
Americium-241 using a CdTe detector;
[0029] FIG. 9 is a graph illustrating the energy spectrum of
Cobalt-57 using a CdZnTe detector;
[0030] FIG. 10 is a flowchart outlining the Monte Carlo gamma ray
history for a modeled system according to the invention;
[0031] FIG. 11 is an illustration of a single head system according
to the present invention;
[0032] FIG. 12 is a schematic diagram of a possible multi-channel
silicon microstrip detector readout chip with fast data readout and
trigger output capability; and
[0033] FIG. 13 is a block diagram of a real time data acquisition
system for use with the present invention.
DESCRIPTION OF THE SPECIFIC EMBODIMENTS
Gamma Ray Detection
[0034] The most probable interaction mechanism for 0.05 to 10 MeV
gamma rays in silicon is the Compton scatter process. Therefore,
the detection of gamma rays in this energy range must use Compton
interaction to have maximum sensitivity. The detector must also
have excellent angular and energy resolution and a wide
field-of-view. The best detection technique that has all of these
features is the Compton double scatter method. This technique
incorporates Compton scattering, photoelectric absorption, and pair
production. The three gamma ray interaction mechanisms are briefly
discussed below.
[0035] Although a number of possible interaction mechanisms are
known for gamma rays in matter, only three major types play an
important role in radiation detection: photoelectric absorption,
Compton scattering, and pair production. Of these, only the first
two play a major roll in emission imaging. All of these processes
lead to the partial or complete transfer of the photon energy to
electron energy. They result in sudden and abrupt changes in the
photon history where the photon disappears entirely or is scattered
through a significant angle.
[0036] FIG. 1 is an illustration of the gamma ray linear
attenuation coefficients for silicon microstrip detectors through
these three processes. The photoelectric absorption dominates below
about 50 keV for silicon. Compton scattering becomes important at
about 50 keV and it stays the dominant process up to about 10 MeV,
where pair production takes over. In the range of 81 to 511 keV
which includes the nuclear medicine range, the important detection
process for silicon is Compton scattering. Compton scattered gamma
ray photons with energies less than 50 keV are readily absorbed due
to the photoelectric effect.
[0037] In the photoelectric absorption process, a photon undergoes
an interaction with an absorber atom in which the photon completely
disappears. In its place, an energetic photoelectron is ejected by
the atom from one of its bound shells. The interaction is with the
atom as a whole and can not take place with free electrons. The
photoelectron appears with an energy, E.sub.c, given by
E.sub.e=h.upsilon.-E.sub.b
[0038] where h.upsilon. is the incident photon energy and E.sub.b
represents the binding energy of the photoelectron in its original
shell. For gamma ray energies, h.upsilon., of more than a 100 keV,
the photoelectron carries off most of the original photon energy.
For silicon microstrip detectors, this process is only important
for low energy gamma rays in the range of 0.5 to 50 keV.
Photoelectric absorption falls nearly exponentially with an
increase in energy. Since the incident photon is totally absorbed
it is not possible to determine the direction of the incident
photon. Therefore collimators must be used to determine the
direction of origin of the photon.
[0039] Compton scattering takes place between the incident gamma
ray and an electron in the absorbing material. In Compton
scattering, the incident gamma ray is deflected through an angle
.theta. with respect to its original direction as illustrated in
FIG. 2. The photon transfers a portion of its energy to the recoil
electron that was initially at rest. Because all angles of
scattering are possible, the energy transferred to the electron can
vary from zero to a large fraction of the gamma ray energy. This
has been a problem in the detection of gamma rays at energies
dominated by the Compton scatter process, since the detected recoil
electron alone does not give sufficient information to uniquely
determine the energy and direction of the incident photon. This has
been solved by the Compton double scatter technique described below
and illustrated in FIG. 2.
[0040] The total incident gamma ray energy, E.sub..gamma., and
Compton scatter angle, .theta., for the double scatter process are
given by:
E.sub..gamma.=E.sub.e1+E.sub..gamma.1
[0041] and
Cos.theta.=1-mc.sup.2(1/E.sub..gamma.1-E.sub..gamma.)
[0042] The incident gamma ray first scatters by the Compton process
in one of the silicon strip detectors 201, losing recoil energy
E.sub.e1. The scattered photon continues on until it interacts with
another silicon strip detector or is absorbed by a calorimeter 203.
If the second interaction is photoelectric absorption, the full
energy of the scatter photon is measured and the energy of the
incident photon and the scatter angle are determined. This is the
dominant process for the calorimeter as it is made of high Z
material and photoelectric absorption increases exponentially with
a decrease in the scattered photon energy. Another possibility is
that the second interaction can be another Compton scatter where
the photon escapes with a small amount of the energy. If the energy
of the escaping photon is sufficiently low, the energy
determination is not significantly effected. If there are enough
silicon planes, the escaped photon makes further interactions in
subsequent planes and gets fully absorbed by the photoelectric
effect. All of the energy measured after the second scatter is just
added to the energy of the second scatter, E.sub.e2, to correct for
the missing energy. If not enough silicon planes are used, for
example due to cost considerations, a calorimeter can be placed
such that it surrounds the sides and the bottom of the silicon
strip detector hodoscope. The surrounding calorimeter is used as a
second scatterer to measure the energy and direction of the
scattered photon or to catch the escaping photons and correct
E.sub.e2 for accurate incident photon scatter angle determination.
Since the calorimeter is a high Z and high density detector or
scintillator, there is a high probability that the escaped low
energy photon will be fully absorbed. The events that do not add up
to the full energy of the incident photon are rejected to reduce
scattered photon background.
[0043] The incident gamma ray direction lies on a cone segment in
the field-of-view with a half-angle .theta.. The cone axis is
determined by the interaction positions in the first and the second
scatters. This is because the direction of the scattered electron
in the top scintillator is not measured. The Compton scattered
electrons with energies in the range of 81 to 364 keV are fully
stopped within 0.03 and 0.3 millimeters of the silicon strip
detectors, respectively. Therefore silicon strip detectors with a
thickness of 0.3 to 2 millimeter are ideal for the present
system.
Silicon Microstrip Detectors
[0044] In the preferred embodiment of the invention, silicon
microstrip detectors are used as the first scatterer (i.e.,
hodoscope). Silicon microstrip detectors have large active areas,
excellent energy and position resolution, and fast readout. Three
inch diameter wafers, typically 200 to 500 micrometers thick, with
parallel readout strips of greater than 25 micrometers pitch on one
side have been available for few years. Pitch size can have any
value from 25 micrometers to several centimeters.
[0045] On the average, 1 electron-hole pair is produced per 3.6 eV
of deposited energy. The energy deposited by an 80 keV recoil
electron fully stopped in silicon is about 22,000 electrons (and
holes) which can be collected in less than 10 nanoseconds. This
leads to pulse rise times of less than 10 nanoseconds. Spatial
resolutions of less than 10 micrometers in one dimension are
obtainable by exploiting charge division between adjacent strips.
Superimposed on the signal is Gaussian-distributed noise related to
the detector strip and preamplifier input capacitances. This noise
or equivalent noise charge (e.g., ENC) is typically about 1,000
electrons at room temperature for detector capacitances of about 20
pF. Thus large signal-to-noise ratios, on the order of 22, are
obtainable for 80 keV electrons.
[0046] To date, silicon detectors have been primarily used in high
energy physics experiments to detect minimum ionizing high energy
charged particles. The Compton converter in the present invention
is different in that the recoil electron loses its entire energy in
a single detector wafer of about 1 millimeter thickness instead of
depositing only part of its energy like the minimum ionizing
particles. The energy and angular resolutions improve as the number
of electron-hole pairs created in the silicon increase. For a 300
keV recoil electron stopping in silicon, about 83,000 electrons
(i.e., 278 e/keV) are produced with an inherent energy resolution
of 0.8 percent (i.e., FWHM/E.sub.0=2.35/{square root}N where N is
the number of electron-hole pairs). For 141 keV electrons stopping
inside the silicon wafer, the theoretical energy resolution is
calculated to be about 1.2 percent with a stopping distance for the
recoil electron of about 0.1 millimeters. The theoretical
resolution can be approached if the input capacitance and the
preamplifier noise can be kept low. The input capacitance can be
decreased substantially by mounting the chips next to the strips or
building them on the same silicon. In the present invention a low
noise, 64 channel front end mixed signal application specific
integrated circuit (ASIC) readout chips is used.
[0047] The individual detector thicknesses can be increased in
order to decrease the number of required planes. Silicon strip
detectors with a 1 millimeter thickness are readily available while
detectors with thicknesses of 2 millimeters have been manufactured.
The energy resolution of silicon strip detectors is a dramatic
improvement over scintillators (e.g., BC-523: 17% at 0.5 MeV).
[0048] Double sided readout silicon microstrip detectors with
orthogonal strips on opposite sides have been developed. FIGS. 3
and 4 show the basic features of a double sided silicon microstrip
or strip detector. The distinct advantage with this configuration
is that both x and y coordinates of a traversing particle are
determined in a single detector plane. For single sided detectors,
the junction side of a standard p+n diode is segmented into many
strips. For double sided detectors, the ohmic side of the n-type
silicon wafer is also segmented with orthogonal strips to provide
simultaneous readout of the particle impact point. Position
resolutions well below a square millimeter on both sides can be
achieved. The preferred detector in the present invention uses 200
to 300 micrometer thick, double sided, silicon microstrip detectors
with about a millimeter spaced strips orthogonal on the top and
bottom surfaces. Such detectors are now commercially available and
fit well with the present design. The x and y positions of the
first two interaction points on the recoil electron track determine
the electron direction. A combination of all interactions is used
to determine the energy of the recoil electron as well as the
scatter angle.
[0049] In one embodiment of the invention, the detector is 6.4
centimeters by 6.4 centimeters, the detector being fabricated from
a 4 inch wafer. In another embodiment, 10 centimeter by 10
centimeter detectors are used. Bridged detectors with overall
lengths exceeding 25 centimeters can also be used with the present
invention. Bridging allows one preamplifier to be connected to a
series of strips on adjacent detectors with significant savings in
electronics.
[0050] A simple Monte Carlo calculation using Monte Carlo Neutron
Photon (MCNP) software from Los Alamos National Laboratory was
performed. The MCNP software gives the probability for a 141 keV
photon to Compton scatter in varying total silicon thicknesses. For
example, about 50 percent of the 141 keV photons will Compton
scatter in a silicon detector 2 centimeters thick. If 2 millimeter
thick silicon strip detectors are used, then 10 planes will be
required. For lower energy photons, a lower total thickness is
required.
[0051] Another important advantage of silicon microstrip detectors
is that they do not need high voltages or cooling to low
temperatures. Room temperature functionality is important to
produce small size, low cost, and low power detectors. They also
have a strong potential for mass production. However, a significant
number of wafers are needed to achieve the conversion rates
required for high sensitivity. Their small thickness and ultrasonic
wire bonding capability render them good candidates for compact
printed circuit board mounting with data acquisition ICs placed
next to them. The readout ICs are preferably designed to give fast
trigger outputs when events occur and output the address and the
analog content of the channel that has the data.
Calorimeter
[0052] Preferably a calorimeter is placed around and at the bottom
of the silicon microstrip detectors in order to absorb the escaping
Compton scattered photons. A variety of different high density
radiation detectors can be used. Many of these detectors are
relatively high cost (e.g., HPGe, BGO, CdWO.sub.4 and CsF) and
several require cooling to liquid nitrogen temperatures (e.g.,
HPGe).
[0053] Sodium Iodide is the most popular high density scintillator.
It has a large light yield and its response to electrons and gamma
rays is close to linear over most of the significant energy range.
The NaI(Tl) crystal is fragile and hygroscopic and therefore must
be handled carefully and hermetically sealed. It has long decay
time and is not suitable for fast timing applications.
[0054] Cesium Iodide is another alkali halide that has gained
substantial popularity as a scintillator material. It is
commercially available with either thallium or sodium as the
activator material and has significantly different scintillation
properties with thallium. CsI has a larger gamma ray absorption
coefficient per unit size and is less brittle than NaI. The two
forms of CsI scintillators, CsI(Na) and CsI(Tl), are discussed
separately below.
[0055] CsI(Na) has an emission spectrum similar to NaI(Tl). Its
light yield is also comparable. CsI(Na) is hygroscopic and must be
hermetically sealed. Therefore, CsI(Na) is similar to NaI(Tl) and
has the same draw backs.
[0056] CsI(Tl) is different than NaI(Tl) and has unique properties.
It is also only slightly hygroscopic. Energy resolution of 5
percent FWHM at 0.662 MeV has been obtained with 2.5 centimeter
diameter by 2.5 centimeter thick CsI(Tl) scintillation crystals
coupled to large area (e.g., 2.5 centimeter diameter) mercuric
iodide photodetectors. This is about 50 percent better than the
NaI(Tl) detectors. The mercuric iodide photodiodes are not yet
available as commercial devices. Resolution of 6 percent at 0.662
MeV has been obtained for considerably smaller CsI(Tl) crystals
using avalanche photodiodes. Large area PIN diodes coupled to 1
centimeter by 2 centimeter CsI(Tl) crystals give a 7 percent
resolution at 0.662 MeV. These crystals produce 35 percent more
photons per MeV than NaI(Tl) and their light spectrum is much
better matched to the sensitivities of the photodiodes. A key to
improved energy resolution is good light collection by matching the
areas of the crystals to those of the photodiodes.
[0057] An important property of CsI(Tl) is its variable decay time
for different particles. Therefore pulse shape discrimination
techniques can be used to differentiate among various types of
radiation such as electrons, protons and alpha particles. The
CsI(Tl) light output is quoted lower than NaI(Tl) for bialkali
photomultiplier tubes (PMTs) (FIG. 5). The scintillation yield is
actually found to be larger than that of any other scintillator
because its light emission peaks at longer wavelengths. It can be
used with photodiodes with extended response in the red region of
the spectrum. Its energy resolution is equal to or better than the
energy resolution of the NaI(Tl) crystals. For these reasons
CsI(Tl) crystals are used in at least one embodiment of the
invention.
[0058] CdTe, CdZnTe, HPGe and HgI.sub.2 are solid state detectors
and can be made in arrays for position sensitive applications. They
are high Z and high density crystals. They are used to detect
x-rays and gamma rays directly without need for photomultiplier
tubes or PIN and avalanche photodiodes. They produce much better
energy resolution than the other detectors which require
photomultiplier tubes or PIN and avalanche photodiodes since they
convert the energy deposited by the x-ray and gamma ray photons
into light, not electron-hole pairs.
[0059] High purity germanium (HPGe) offers excellent high energy
resolution and exhibits moderate gamma ray absorption properties,
making it the detector of choice for high accuracy spectroscopy.
Unfortunately since it only works at liquid nitrogen temperatures,
bulky refrigeration systems are required which further increase the
cost of this detector. HPGe is available in single small crystals
and works by collecting the electron hole pairs produced inside the
crystal similar to the silicon detectors and does not require PMTs.
Because of the large cost this detector is mainly used for
applications which require ultra high energy resolution and small
size detectors.
[0060] BGO, CdWO.sub.4 and CsF are excellent high density and high
Z scintillators. They have lower energy resolution and light
output. Their maximum light emissions peak around 430 nanometers,
similar to NaI(Tl), and require PMTs for detection. CdWO.sub.4 and
especially CsF have shorter decay constants and faster rise times
than the others and can be used for timing. However, since the
preferred detector of the present invention does not use
time-of-flight to determine the direction of the scattered gamma
ray photon, good time resolution is not important.
[0061] The preferred room temperature detector for the calorimeter
of the present invention is CdTe or CdZnTe. These detectors are
described in more detail below.
System
[0062] The present invention, relying on isotope uptake in the
region (i.e., organ or tissue) of interest, can be used for a
variety of different applications ranging from real-time monitoring
(e.g., blood flow through a heart valve) to lesion diagnosis (e.g.,
breast lesions). The disclosed system is relatively compact while
offering improved efficiency and spatial resolution. An obvious
benefit of the improved efficiency of the present invention is a
significant decrease in the observation time or the
radiopharmaceutical dosage.
[0063] One embodiment of the invention is illustrated in FIGS. 5
and 6. An object 501 to be imaged such as a breast, brain, or other
organ or tissue is placed at the front of system 500. The hodoscope
is made up of between approximately 1 and 100, and preferably
between approximately 10 and 25, silicon strip detector planes 503.
Detector planes 503 preferably have a thickness of between 0.5 to 1
millimeter, the selected thickness being dependent upon the desired
performance as well as the availability of the detectors. The total
Compton scatter probability will vary from approximately 50 percent
for ten 2 millimeter thick silicon strip detectors to approximately
35 percent for twenty five 0.5 millimeter thick detectors. The
active area of the silicon strip detectors can be increased by
mounting four or more detectors 601 side by side on each plane 503
as shown in FIG. 6.
[0064] The hodoscope height depends strongly on the number of
detector planes 503 as well as the separation between the planes.
In the illustrated embodiment, 1 millimeter thick detectors are
used giving a plane separation of about 1 centimeter and a
hodoscope height of about 15 centimeters. Preferably these values
as well as the thickness and separation of the silicon detectors is
optimized through Monte Carlo simulations and experimental
study.
[0065] The calorimeter is made from about 2 millimeter thick CdTe
or CdZnTe strip or pad detectors 505. The reason for this selection
is the higher energy resolution obtained from CdTe/CdZnTe detectors
especially at lower energies. A CsI(Tl) calorimeter can also be
used.
[0066] In the illustrated embodiment, calorimeter 505 is a single
layer placed around, and as close as possible, to the hodoscope.
The proximity of the calorimeter to the hodoscope is limited in
order to avoid introducing significant angular resolution
degradation due to the geometric combination of pixels. The gap at
the bottom is due to the energy threshold of the silicon detectors
which is typically greater than 5 keV. The incident photons that
deposit energy less than the threshold energy will not be detected
in the hodoscope and such small angle scatters need not be stopped
at the calorimeter. The geometry, strip pitch, thickness,
shielding, and the size of the gap at the bottom of the calorimeter
is optimized by Monte Carlo simulations. The detector geometry is
optimized to any form such as square, rectangular, cylindrical,
spherical, parabolic, etc. that gives the best results for a
specific application.
[0067] A shield 507, preferably made of a material such as lead or
tungsten, is placed in front of and around the calorimeter to
reduce the background. Shield 507 is especially important for
certain applications. For example, if imaging system 500 is used
-15 for the detection of malignant breast tumors, a
radiopharmaceutical such as Tc-99m SestaMIBI or Tl-201 may be used.
In either case, substantial amounts of the radiopharmaceutical may
be taken in by the heart thus requiring adequate shielding to
achieve the desired signal-to-noise ratio.
[0068] In an alternate embodiment used to obtain an intermediate
improvement in sensitivity, a slot collimator 509 is placed at the
aperture of the hodoscope. Collimator 509 confines photons to the
planes defined by the slots. Collimator 509 will therefore slice
the event cone inherent in a Compton scatter detector into two
sections, one section defining the true event direction and the
other section defining the false event direction. The false event
directions normally lie outside the viewed object, especially for
large scatter angles. Thus the correct and incorrect directions can
be defined for each event and all false events can be rejected.
[0069] In the preferred embodiment of the invention, each plane of
the hodoscope is made from four 1 millimeter thick silicon strip
detectors 601 with an active area of approximately 6.4 centimeters
by 6.4 centimeters each. Detectors 601 are mounted as close to each
other as possible. Therefore in this embodiment the active area is
approximately 12.8 centimeters by 12.8 centimeters or about 164
square centimeters. The number of detector planes 503 is a function
of the application. For example, if system 500 is to be used to
image breast tumors, the hodoscope preferably has 10 detector
planes with a 0.5 centimeter spacing. For other organ imaging
applications the hodoscope has between 15 and 25 detector planes
with an approximately 1 centimeter spacing.
[0070] Silicon strip detectors 601 are mounted on a printed circuit
board (PCB) 603 or a ceramic holder as illustrated in FIGS. 6 and
7. Front end electronics (FEE) readout chips are mounted on the PCB
proximate to the silicon strip detectors 601 either on the front
surface of the PCB at locations 701 or on the back surface of the
PCB at locations 703. A fan-in from the strip pitch to the FEE chip
pad pitch is done on the silicon strip detector for reliability and
ease of ultrasonic wire bonding.
[0071] Preferably silicon strip detectors 601 are designed and
fabricated using the new FOXFET AC coupling technique on both the
junction and ohmic sides. This technique improves the signal
quality, especially at the ohmic side, since the bias resistor
formed through the FOXFET technique is much larger than with other
techniques. It also eliminates external capacitances and resistors
which become bulky, require large real estate, and are costly when
large numbers of channels are used. Preferably high radiation
resistant FOXFET silicon strip detectors are used which
significantly increases the reliability of the system.
[0072] FOXFET silicon strip detectors are commercially available
and show excellent response to low energy (i.e., 81 to 511 keV)
photons. By lowering the dark current and reducing the junction
thickness to decrease strip capacitance, a reduction in detector
and electronic noise should be achieved, thereby improving energy
and spatial resolution.
[0073] The small size of the active area and the dividers between
the four silicon strip detectors at each plane do not cause
problems such as side truncation or image gaps. This is because the
disclosed technique inherently has a large field of view and the
detector active area can be smaller than the imaged organ of the
patient. Smaller active area, dead strips within a plane, or gaps
in between the silicon strip detectors, only reduce the detection
efficiency while not affecting the image. Thus as a result of the
Compton scatter technique, image defects are virtually eliminated
due to the system's tolerance to defects.
[0074] Although the invention can be used without a calorimeter,
the preferred embodiment includes a calorimeter utilizing CdZnTe
strip detectors. These detectors have excellent energy resolution
for 10 to 250 keV gamma rays, or for 250 to 600 keV gamma rays if
thick detectors are used. Therefore, CdZnTe is especially useful to
work with .sup.99mTc and .sup.201Tl, the most commonly used
radionuclides.
[0075] The second choice for the calorimeter are CsI(Tl) crystals
coupled to specially developed PIN photodiodes. The energy
resolution of these crystals, contrary to CdTe detectors, increases
as the gamma ray energy increases. Therefore, they are an excellent
choice for source gamma rays with energies greater than 250
keV.
[0076] At higher energies, the thickness of silicon required to
stop the gamma rays becomes larger, requiring multiple Compton
scatters prior to absorption. If a calorimeter is used, the
incident photon only needs to make a single Compton scatter in the
silicon hodoscope.
[0077] The energy resolution for a 1 by 1 by 2 cubic centimeter
crystal of CsI(Tl) is approximately 5 percent at 662 keV using a
.sup.137Cs source, thus showing that a CsI(Tl) calorimeter can be
used with the present invention. A CsI(Tl) calorimeter with smaller
crystals can be used at lower energies without a significant impact
on the stopping power of the calorimeter. For example, a 0.5
centimeter long CsI(Tl) crystal can absorb 95 percent of 141 keV
photons.
[0078] In general the present invention has three basic embodiments
depending upon the intended use. The first embodiment is intended
for relatively low energy, i.e., between about 81 and about 250
keV. Due to the low energy, this embodiment can be fabricated
either with or without a calorimeter. If a calorimeter is used,
preferably it is a CdZnTe calorimeter. The second embodiment is
intended for relatively high energy, i.e., between about 250 and
about 511 keV or greater. This embodiment uses both the hodoscope
and the calorimeter, the calorimeter utilizing either CdZnTe or
CdI(Tl). The third embodiment can be used throughout the entire
energy range, albeit with slightly lower efficiency and spatial
resolution. In this embodiment preferably a CsI(Tl) or a thick
(e.g., 0.5 to 1 centimeter) CdZnTe calorimeter is placed behind a 2
millimeter thick CdZnTe plane. The CdZnTe calorimeter is useful for
low energy radionuclides while both the CdZnTe and the CsI(Tl)
calorimeter can be used with the silicon hodoscope for high energy
sources. In such an arrangement interactions in all three sections
may happen and can be used as viable data for imaging.
[0079] The origin of CdZnTe is the cadmium telluride (CdTe)
detector. CdTe contains relatively high atomic numbers (i.e., 48
and 52) with a large enough bandgap energy (i.e., 1.47 eV) to
permit room temperature operation. This bandgap limits
resistivities to the low-109 ohm-centimeter range, resulting in
relatively large room temperature dark currents. CdTe has a density
of 6.06 grams per cubic centimeter and the energy required to
create a single electron-hole pair is 4.43 eV. The hole mobility is
about a factor of 30 slower than the electron mobility. The hole
life times are also very short due to the low mobility enhancing
the effects of trapping and recombination. Improvements in hole
collection efficiency can be obtained by using higher purity
materials.
[0080] In CdTe, for typical gamma ray energies the probability of
photoelectric absorption per unit pathlength is approximately 100
times larger than in silicon. For example, CdTe is opaque to low
energy x-rays for thicknesses in the range of a millimeter.
However, the energy resolution of CdTe is not comparable to silicon
detectors for low energy x-rays due to poor hole collection
efficiency. The room temperature measured energy resolution for
CdTe detectors is 3.5 keV at 122 keV.
[0081] Many problems associated with CdTe detectors are related to
a specific technique of crystal growth referred to as the traveling
heater method (THM). This technique requires that the crystals be
doped with an element such as chlorine in order to achieve high
resistivity. Unfortunately, chlorine doping is generally associated
with detector long-term reliability problems as well as various
operating instabilities such as counting rate polarization. Lastly,
due to the low yield of detector grade material using this
technique, detector prices are relatively high.
[0082] The CdZnTe detectors were specifically developed as gamma
ray detectors by several companies. By using a high pressure
Bridgman (HPB) technique to grow the crystals, improvements in both
size (e.g., up to 10 centimeter diameter crystals weighing over 10
kilograms) and yield (e.g., over 70 percent) have been realized.
These crystals exhibit uniform, near-intrinsic resistivity without
doping. Detectors fabricated from HPB grown crystals exhibit
excellent stability, reliability and lifetime. Furthermore, the HPB
process can be used to grow high quality crystals of
Cd.sub.1-xZn.sub.xTe throughout the entire alloy composition range.
Alloying ZnTe with CdTe increases the bandgap, resulting in much
higher resistivities and correspondingly lower leakage currents
than CdTe.
[0083] The energy resolution of both the CdTe and CdZnTe detectors
for 10 to 300 keV energies is important for the present invention.
FIG. 8 shows the energy spectrum of an Americium-241 source with a
CdZnTe detector. The x-ray emissions at 13.9, 17.7, 20.8, 26.4, and
59.5 keV (with escape peaks for characteristic K x-rays from Cd at
36.5 keV and Te at 32.5 keV) are clearly seen with good energy
resolution. The slightly lower energy tail observed for the 59.5
keV peak is typical of that observed with CdZnTe detectors and is
due to incomplete charge collection for some of the events.
[0084] The energy spectrum of a Cobolt-57 obtained by a 2
millimeter thick CdZnTe crystal is shown in FIG. 9. The low energy
tail is clearly seen at higher energies.
[0085] In one embodiment of the invention, CdZnTe strip detectors
produced from Cd.sub.0.8Zn.sub.0.2Te wafers were used. The strip
pitch was 1 millimeter with a total of 32 strips on each side,
providing an active area of 3.2 centimeters by 3.2 centimeters. The
strips on each side were orthogonal to each other in order to
provide both the x and y dimensions for an interaction. The CdZnTe
strip detectors can be from 1.5 to 2.5 millimeters thick.
Two-dimensional arrays of CdZnTe pad detectors can also be used.
Pad detectors generally provide better results than strip detectors
since they do not have the positional ambiguity associated with
strip detectors when there is more than one event simultaneously
interacting with multiple detectors.
[0086] A full size cylindrical system according to the present
invention was modeled using the MCNP Monte Carlo code discussed
above. The internal and external radii of this cylindrical system
were 15 and 50 centimeters, respectively. The length of the modeled
system was 50 centimeters long. The phantom used at the center was
a standard cylinder 20 centimeters in diameter 20 centimeters long
filled with water and 1 .mu.Ci/cc of a .sup.99mTc radiotracer.
Double-sided silicon strip detectors that were 1 millimeter thick
with a 1 millimeter strip pitch were modeled in cylindrical form.
All together, 36 planes were placed inside the detector with a 1
centimeter separation between planes. The total thickness of the 36
planes is 3.6 centimeters corresponding to a 72 percent Compton
scatter probability. The 141 keV gamma rays, produced uniformly in
all directions in the phantom, were tracked along their paths until
they were fully absorbed or escaped through the back or sides of
the detector. A 3 keV energy threshold of detection was imposed on
each silicon detector. A calorimeter behind or at the sides of the
model was not used.
[0087] The history of the 141 keV photon was traced by Monte Carlo
calculations, the results of which are shown in FIG. 10. The Monte
Carlo calculations were carried out for about 100,000 events and
the results scaled to the simulated phantom. The 141 keV photons
scattered in the phantom are effectively discriminated by the high
energy resolution. This significantly reduces the major scattered
photon background. The single scatter photons are rejected as their
directions cannot be measured. The events which create multiple
electrons in the same detector wafer are also rejected. Most of
these are probably due to knock on electrons by the recoil
electron. In most cases the secondary electrons are created and
absorbed within the pixel size at the position of the interaction.
These events are legitimate and can be used in imaging.
[0088] The forward and backscattered gamma ray events can be easily
identified because of the strict relationship imposed by the
Compton scatter formula. This is especially true at low photon
energies. For 141 keV .sup.99mTc gamma rays, the energies deposited
in the interaction point nearest to the patient are limited to 0 to
31.5 keV and 110.5 to 90.9 keV for forward and back scattered
photons (i.e., .theta..ltoreq.90.degree.), respectively. For the
interaction point farthest from the patient (i.e.,
90.degree..ltoreq..theta..ltoreq.180.deg- ree.), the energies
deposited are 31.5 to 50.1 keV and 141 to 110.5 keV for the forward
and back scattered photons, respectively. Since the energy at each
interaction plane is measured separately, such widely differing
energy deposition for the forward and backward scattered photons is
easily identifiable and the direction cones can be calculated.
Therefore the backscattered events that deposit full energy in the
detector are good events and can be used in imaging.
[0089] The point sensitivity is estimated to be about 1,500 Cts
s.sup.-1 .mu.Ci.sup.-1. The volume sensitivity of the simulated
detector is about 500,000 Cts s.sup.-1 cm.sup.-1 found by dividing
the good event rate, 1.times.10.sup.7 cts s.sup.-1, by the length
of the phantom. The sensitivity of the invention strongly depends
on the amount of silicon used and can be improved further by
increasing the number of silicon detectors. The number of silicon
strip detectors can also be decreased to reduce cost since the
sensitivity is high and some sacrifice is affordable.
[0090] The FWHM uncertainty in the cone half-angle, .DELTA..theta.,
due to a detector of finite energy resolution (FWHM),
.DELTA.E.sub.c1 and .DELTA.E.sub.c2 at first and second scattering
planes can be calculated using the Compton scatter formula: 1 = mc
2 E 2 Sin { E e1 2 + [ E 2 E 1 2 - 1 ] 2 E e2 2 } 1 / 2
[0091] where mc.sup.2 is the electron rest energy (511 keV),
.theta. is the Compton scatter angle, and E.sub..gamma. and
E.sub..gamma.1 are the incident and scattered photon energies.
Applying the formula, the energy resolution due to the statistical
fluctuation for electrons stopped inside the silicon microstrip
detectors varies from 1.3 percent at 100 keV to 0.75 percent at 350
keV. The electronics noise of the detector is about 2 keV.
Therefore the total energy resolution is dominated by the
electronics noise which is the same for both the converter and the
calorimeter.
[0092] The angular resolution is calculated with an energy
resolution of 2 keV (FWHM) where .DELTA..theta. for forward
scattered gamma rays (i.e., .theta.<90.degree.) varies from
5.degree. at a .theta. of 30.degree. to about 3.2.degree. at a
.theta. of 70.degree. for 141 keV (.sup.99mTc) incident photons.
The same calculation carried out for 364 keV .sup.131I gamma rays
gives angular resolutions of approximately 1.degree. for a .theta.
of between 20.degree. and 90.degree.. Thus the angular resolution
improves significantly with an increase in the photon energy. Also
the effects of amplifier noise are reduced as more electron-hole
pairs are created by higher energy scattered electrons. At a
distance of 20 centimeters these angular resolutions produce
effectively 6 to 3.5 millimeter spatial resolutions for 141 keV
gamma rays and 3.5 to 1.5 millimeter spatial resolutions for 364
keV gamma rays. At a distance of 2.5 centimeters the same energy
gamma rays produce 2.2 to 1.4 millimeter spatial resolutions and
0.4 millimeter spatial resolutions, respectively.
[0093] The geometric angular resolution, .DELTA..theta..sub.Geom,
gives the axis of the image cone and is dependent upon the silicon
microstrip detector pixel size and the distance between the first
two scatters. The FWHM value can be calculated similar to that for
a collimator. Normally the geometric angular resolution is kept
much smaller than the scatter angle variation which depends
strongly on the energy resolution as shown above. It is easier to
adjust the geometric angular resolution in a silicon microstrip
detector as the strip or pixel pitch dimensions can be as small as
25 microns. The pixel size for the simulated model is 1 square
millimeter.
[0094] The Monte Carlo analysis shows that about 1.times.10.sup.8
photons per second out of 2.3.times.10.sup.8 enter the detector as
shown in FIG. 10. As noted above, the simulated detector has 36
cylindrical planes with an average area of approximately 10.sup.4
square centimeters and about 75 percent of the photons making an
interaction (i.e., 7.5.times.10.sup.7 photons per second). Assuming
each silicon microstrip detector wafer has dimensions of 5
centimeters by 5 centimeters, the singles rate in each wafer is
about 5,000 Cts/s. Such singles rates are not excessive for silicon
microstrip detectors which produce about 20 nanosecond long pulses.
The coincidence requirement further reduces the actual readout rate
to about 670 per second. Therefore dead time per detector is not a
problem. However, the total count rates of the whole detector will
be high. This problem is solved by establishing high level
parallelism in readout electronics for which the silicon microstrip
detectors are highly suitable. One possible way is to divide the
detector into many radial sections and read each section
individually. If it is divided into 20 sections than readout rate
at each section will be about 500 kHz which can be easily handled
by a standard CAMAC data acquisition system. The data rate will be
even smaller due to some loss of events at the edges when the
photons scatter into adjacent sections. This will also reduce
sensitivity somewhat unless such events can be recovered by the
electronics. There is also a large number of channels to readout.
This is solved by using high density ASIC chips directly connected
to the microstrips. Chips which produce a trigger signal when there
is valid data and connect the strip that contains information to
the output can be used.
[0095] FIG. 11 is an illustration of a simple, single head
apparatus fabricated according to the present invention and used
for system testing. The hodoscope is comprised of 10 layers 1101 of
silicon strip detectors in which each detector layer 1101 has an
area of 12.4 centimeters by 12.4 centimeters with a thickness of 1
millimeter. The distance between each layer 1301 is 0.5
centimeters. A calorimeter 1105 is symmetrically positioned 6
centimeters after the last hodoscope layer 1101. Calorimeter 1105
is a 2 millimeter thick CdTe detector with an area of 50
centimeters by 50 centimeters. A 141 keV gamma ray source 1107 with
a 0.5 centimeter diameter is centrally positioned 10 centimeters
above the first silicon plane 1101. A threshold energy of 10 keV
was applied to the silicon strip detectors. An event is generated
only if the incident gamma ray makes a Compton scatter in one of
the silicon planes and also interacts at the calorimeter.
[0096] There were a total of 56,234 triggers out of 106 incident
gamma rays. The low efficiency, about 5.6 percent, was due to the
overall small silicon thickness of 1 cm (i.e., approximately 30
percent Compton scatter probability). The low efficiency was also a
result of calorimeter 1105 not covering the sides of the hodoscope
since most of the photons scattered at an angle of greater than
70.degree. would not be detected. Lastly, since the thickness of
calorimeter 1105 was only 2 millimeters, the absorption probability
was approximately 85 to 50 percent for 90 to 131 keV scattered
photons due to the Compton geometry. About 0.56 percent of the
incident photons produced a photoelectric absorption inside the
silicon hodoscope. The total number of events that deposited full
energy in the detector was 4.2 percent. If the events are
restricted to total absorption in calorimeter 1305 after Compton
scattering once in the silicon hodoscope, about 2.8 percent of the
incident photons were detected. This excludes totally absorbed
events in silicon after 2 or more Compton scatters in the
hodoscope.
[0097] Two more sources, 1109 and 1111, were added to the above
single source discussed above. Source 1109 was positioned 2
centimeters from the center source in the -x direction while source
1111 was positioned 1.5 centimeters in the +x direction. All the
sources produced 141 keV gamma rays sprayed into a cone the size of
the hodoscope. The photons produced by the sources at the sides
missed part of the detector aperture due to their position.
Therefore, the strongest source imaged was the centrally placed
one. The images were obtained using a standard analysis program.
This program integrated the overlap of each event ring at the
corresponding pixel. The energies deposited at the hodoscope and
the calorimeter are randomly Gaussian distributed using the
calculated energy resolutions for the preferred prototype system to
simulate authentic spatial resolution.
[0098] The present Compton double scatter detectors provide two
basic parameters for each event related to the incident photon
direction; the scattered photon direction and the Compton scatter
angle. The Direct Linear Algebraic Deconvolution (DLAD) technique
can be used to analyze this information.
[0099] A concise explanation of the DLAD technique is provided
below. The reconstruction of the source image from the Compton
double scatter data can be represented by the following general
formula:
D(.chi.,.PSI.,.PHI.,E)=.intg..sub..chi..PSI..PHI.EI(.chi..sub.0,.PSI..sub.-
0,E')R(.chi.,.PSI.,.chi..sub.0,.PHI.,E',E)d.chi..sub.0d.PSI..sub.0dE'+B(.c-
hi.,.PSI.,.PHI.,E)
[0100] In the above formula, D(.chi.,.PSI.,.PHI., E) is the actual
Compton scatter data observed by the detector in appropriate
coordinates; .chi. and .PSI. are the coordinates of the rectangular
image plane; .PHI. is the Compton scatter angle; E is the energy of
the incident photon; I (.chi..sub.0, .PSI..sub.0,E.sub.0) is the
true image of the source and is not a function of the Compton
scatter angle; R(.chi., .PSI., .chi..sub.0, .PSI..sub.0, .PHI., E',
E) is the response function of the detector; and B(.chi., .PSI.,
.PHI., E) is the gamma ray background. Normally the calculation is
carried out for all energies within the detector sensitivity to
determine the total gamma ray flux and for certain energy bands to
obtain an energy spectrum. For application to the present
invention, the energy spectrum is used to discriminate the
scattered photon background. The calculation can also be done for
different scatter angle bands. D and I are normally referred to as
the data and the image spaces, respectively.
[0101] The response function in the DLAD technique is the
concentric rings obtained by mapping the scattered photon direction
vector in the image plane. This can be used as an ideal detector
response function. The true detector response function, R, can be
represented by
R.sub.i
j,.PHI..sub..sub.s=.epsilon.(E,.theta..sub.j,.PHI..sub..sub.s).mul-
tidot..DELTA..PHI..sub.s.multidot.PSF.multidot.G(.theta..sub.i)
[0102] where i and j define the bins in the data and image spaces,
respectively; .PHI..sub.s is the calculated Compton scatter angle
as given by Compton scatter formula; .epsilon. is the detector
efficiency; .theta..sub.i and .theta..sub.j are the incident zenith
angles in data and image spaces, respectively; .DELTA..PHI..sub.s
is the scatter angle interval; PSF is the point spread function;
and G(.theta..sub.i) is the geometric factor. The PSF is the
distribution of the scattered photon vectors in the image plane.
The PSF can be represented by the two dimensional normal
distribution
PSF=C(.theta..sub.j,.PHI..sub.s)e.sup.-{[(.PHI..sup..sub.t.sup.-.PHI..sup.-
.sub.s.sup.).sup..sup.2.sup.]/[2.sigma..sup..sup.2.sup.(E)]}
[0103] where C is the normalization constant determined by the
requirement that PSF.times.G(.theta..sub.i) is equal to 1. The PSF
and G(.theta..sub.i) are symmetric in the azimuth, thus giving a
two-dimensional image. The present invention can produce
three-dimensional images due to the Compton scatter process.
Therefore, either two-dimensional image slices parallel to the
converter planes are produced or a direct three-dimensional image
can be constructed.
[0104] The DLAD technique can produce fluctuations on the image
space that are due to the geometric factor forcing data space to
zero at the corners and edges of the field-of-view where the data
may be scarce and the Poisson fluctuations are large. This effect
can be improved by applying the positivity requirement. The
positivity requirement is based on the fact that in image space one
cannot get negative fluxes. The positivity constraint has been
introduced into DLAD. The new technique is called Constrained
Linear Algebraic Deconvolution (CLAD).
[0105] An important technological requirement for the present
invention is a multichannel front end electronics (FEE) chip with
self trigger output to readout the silicon strip and calorimeter
detectors. A detailed description of a FEE chip is provided in U.S.
Pat. No. 5,696,458, issued Dec. 9, 1997 and in co-pending U.S.
patent application Ser. No. 08/866,117, filed Jun. 27, 1997, both
disclosures of which are incorporated herein for all purposes.
[0106] The preferred FEE chip is a 64 channel, charge sensitive,
mixed signal ASIC CMOS chip, a version of which is illustrated in
FIG. 12. Each channel of the chip consists of an analog section and
a digital section. The input from the silicon strip detector is
directly coupled to a low noise, charge sensitive amplifier. The
outputs of the charge sensitive amplifier are connected to a shaper
amplifier with a time constant of about 100 to 200 nanoseconds. The
output of the shaper amplifier goes into the track and hold (T/H)
switch. The T/H switch can be controlled externally or activated
internally from the trigger output with a delay set to turn on the
hold at the peak of the shaped pulse. The T/H switch is connected
to the input of the buffer amplifier through the voltage following
capacitor. When the T/H switch is open the voltage on the capacitor
is held constant and the voltage level is buffered on to the analog
output switch. A shift register connects each buffer output to the
single analog output pin in sequence, from input 1 to N, by an
external clock input. The shift register also has an external clear
input to reset it and a clock output to daisy chain it to other
readout chips. Only one clock input is sufficient if the clock
outputs are connected in serial to the clock inputs of the adjacent
readout chips. The charge sensitive amplifier outputs can be fanned
out to comparators with a common external level adjustment. The
outputs of the comparators can be fanned in through a fast OR
circuit which will produce a trigger signal if any comparator input
exceeds the set threshold. The trigger signal can also be used with
a suitable delay to control the T/H switches to apply hold signal
at the peak of the pulse from the shaper amplifier.
[0107] The data acquisition speed of the readout chip will also be
increased using the extra versatility introduced by the
comparators. The design shown in FIG. 12 does not tell which strip
has the information so all strips are readout to find the strip
that has the signal. A logic circuit can be added to the design
which detects the channel with the largest signal from the
comparator outputs, applies a track and hold signal, and connects
the strip with the signal to the analog output pin. At the same
time it can encode the address of the strip that has the
information and output it as the address of the strip with the
signal. There could be an occasional signal on more than one strip.
Multi-hits can be detected and an output can be generated to warn
of a multi-hit signal. The trigger signals are generated for each
readout chip. They have to be externally processed for the
hodoscope in coincidence with the calorimeter to produce the single
trigger signal to activate the data acquisition system. For
extremely high signal rates this may not be possible. In such a
case each wafer or front end readout chip can be separately readout
in parallel using independent data acquisition electronics and
tagging each event time by using an accurate clock. The calorimeter
crystals are also individually readout and event times tagged by
the same clock. Since the calorimeter is running at much slower
speeds, individual readout modules are not necessary and can be
readout in groups.
[0108] The data readout can be carried out in parallel and can be
stored on-board using individual module memory. This is the key to
achieve fast data throughput rates. The data can be asynchronously
accessed by the host computer, analyzed and displayed on screen in
real time. Data acquisition rates of 1 to 10 MHz per readout chip
(or silicon wafer) are achievable.
[0109] A block diagram of the readout electronics system is shown
in FIG. 13. The electronics has two similar sections for the
hodoscope and the calorimeter readout. A true event is a
coincidence between the hodoscope and the calorimeter. The two
master trigger signals from the hodoscope and the calorimeter are
sent to a coincidence unit to create the Compton double scatter
event trigger. The Compton double scatter trigger signal is only
generated if there is a master trigger signal from both the
hodoscope and the calorimeter. This is the arrangement which does
not employ the time tagged data readout method. Time tagged data
acquisition will only be used if absolutely necessary.
[0110] The Compton double scatter event trigger activates data
acquisition for both the hodoscope and the calorimeter
simultaneously. Either CAMAC or VME bus modules can carry out the
data acquisition. The CAMAC system is the most cost effective.
Faster computer interface busses such as Fastbus, VME or VXI bus
can also be used. The custom designed data acquisition modules for
the hodoscope will produce the necessary microstrip readout chip
control electronics, such as the T/H (if not generated internally
in the readout chip), a clear signal to reset the shift registers,
and the clock pulse to multiplex each strip to the analog
output.
[0111] The analog input channels from different hodoscope planes
are read out synchronously with the clock pulse output. The module
converts the pulse height information received from the analog
output pin to a digital number. In parallel with reading the
hodoscope data, it also digitizes the signal(s) from the
calorimeter. Immediately after reading out the last signal it
clears the hodoscope to reset the readout chip so that it can
receive the next event. It is assumed that the analog output of
each readout chip in each detector plane is fanned in to allow a
single signal to be sent to the readout module. It is also possible
to design a microstrip readout chip that can internally connect the
strip which has the maximum signal to the analog output and also
produce the encoded address of the strip. In such a case the clock
output will not be necessary and the silicon microstrip detectors
can be readout asynchronously at a much faster rate.
[0112] The custom made CAMAC modules are connected to the CAMAC
crate controllers which are standard devices and available off the
shelf. The controllers connect the modules to the data acquisition
computer. Depending on the data rate and readout overhead, single
or separate computers can be used to read the hodoscope and the
calorimeter. The computer stores data on a hard disk, optical
drive, or nonvolatile RAM depending on the application. If the data
acquisition overhead is not high then one of the computers can
analyze the data in real time or a separate computer can access the
storage media asynchronously. The results of the data analysis are
imaged onto the field-of-view through a display system in real
time.
[0113] The data analysis techniques for nondestructive evaluation
inspection resemble closely those of medical Computer Assisted
Tomography (CAT) imaging. This type of imaging is based on the
Radon transform and back projection techniques and is standard in
the industry. New iterative techniques such as Maximum Likelihood
and Maximum Entropy methods can also be applied to enhance the
image quality as can the DLAD technique described above.
[0114] If a calorimeter is not used the direction and the energy of
the incident photon has to be measured in the hodoscope. This can
be achieved by increasing the total thickness. These measurements
can be made by two scatters where the second scatter is a
photoelectric absorption. If an incident photon makes 3 or more
scatters (i.e., it is over determined), then the Compton scatter
angle and the energy of the incident photon can be determined more
than one independent way even if the photon does not deposit its
full energy in the silicon converter and escape. Such multiple
Compton scatters can also lead to a reduction in the azimuthal
ambiguity (i.e., event ring) because the Compton scattered photon
will be polarized and the third interaction position is dependent
on the scattered photon direction.
[0115] Monte Carlo simulation of a hodoscope only system has been
carried out using the same geometry as discussed above without a
CdZnTe calorimeter and with 20 silicon strip detector planes. Out
of 3.times.10.sup.6 141 keV gamma rays incident on the detector, 22
percent made a single scatter and escaped out, 10 percent of which
were absorbed by the photoelectric effect. 7.7 percent of the
incident gamma rays made two scatters, 20 percent of these
depositing their full energy in the hodoscope. 4.2 percent of the
incident gamma rays produced 3 or more scatters. Most of these, in
theory, may be used to determine the incident photon energy and
scatter angle.
[0116] If a pure single line source is used then a high sensitivity
imaging mode can be applied with some reduction in spatial
resolution. In this mode the background discrimination cannot be
applied for double scatters. The requirement that the double
scattered photon must deposit all of its energy in the hodoscope
reduces the number of useful events for imaging by 80 percent.
Since the energy of the incident photon is known than the missing
energy of the escaped photon can be added to the second scatter and
the scatter angle can be determined. This method improves the
signal somewhat but also increases the background. However, for a
hodoscope only system it may increase the good data rate by a
significant factor.
[0117] As will be understood by those familiar with the art, the
present invention may be embodied in other specific forms without
departing from the spirit or essential characteristics thereof.
Accordingly, the disclosures and descriptions herein are intended
to be illustrative, but not limiting, of the scope of the invention
which is set forth in the following claims.
* * * * *