U.S. patent application number 14/536956 was filed with the patent office on 2015-10-08 for position sensitive solid-state photomultipliers, systems and methods.
This patent application is currently assigned to Radiation Monitoring Devices, Inc.. The applicant listed for this patent is Radiation Monitoring Devices, Inc.. Invention is credited to James F. Christian, Purushottam Dokhale, Mickel McClish, Kanai S. Shah, Christopher Stapels.
Application Number | 20150285921 14/536956 |
Document ID | / |
Family ID | 51845761 |
Filed Date | 2015-10-08 |
United States Patent
Application |
20150285921 |
Kind Code |
A1 |
Shah; Kanai S. ; et
al. |
October 8, 2015 |
POSITION SENSITIVE SOLID-STATE PHOTOMULTIPLIERS, SYSTEMS AND
METHODS
Abstract
An integrated silicon solid state photomultiplier (SSPM) device
includes a pixel unit including an array of more than 2.times.2 p-n
photodiodes on a common substrate, a signal division network
electrically connected to each photodiode, where the signal
division network includes four output connections, a signal output
measurement unit, a processing unit configured to identify the
photodiode generating a signal or a center of mass of photodiodes
generating a signal, and a global receiving unit.
Inventors: |
Shah; Kanai S.; (Watertown,
MA) ; Christian; James F.; (Waltham, MA) ;
Stapels; Christopher; (Mills, MA) ; Dokhale;
Purushottam; (Waltham, MA) ; McClish; Mickel;
(Salem, MA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Radiation Monitoring Devices, Inc. |
Watertown |
MA |
US |
|
|
Assignee: |
Radiation Monitoring Devices,
Inc.
Watertown
MA
|
Family ID: |
51845761 |
Appl. No.: |
14/536956 |
Filed: |
November 10, 2014 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
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12580172 |
Oct 15, 2009 |
8884240 |
|
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14536956 |
|
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61185169 |
Jun 8, 2009 |
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Current U.S.
Class: |
600/411 ;
250/361R; 250/370.14 |
Current CPC
Class: |
G01T 1/208 20130101;
G01T 1/1603 20130101; G01T 1/249 20130101; G01T 1/18 20130101; G01T
1/248 20130101; G01T 1/2985 20130101; G01R 33/481 20130101; G01T
1/2018 20130101 |
International
Class: |
G01T 1/18 20060101
G01T001/18; G01T 1/24 20060101 G01T001/24; G01R 33/48 20060101
G01R033/48 |
Goverment Interests
STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT
[0002] This invention was made with government support under grant
number DE-FG02-08ER84988 awarded by the Department of Energy and
grant number 2R44NS060197-02 awarded by the Department of Health
and Human Services. The government may have certain rights in this
invention.
Claims
1. An integrated silicon solid state photomultiplier (SSPM) device
operable in a Geiger mode, the device comprising: a pixel unit
including an array of more than 2.times.2 p-n micro-pixels on a
common substrate; a signal division network electrically coupled to
each micro-pixel, the signal division network including four output
connections each for providing an output signal; a signal output
measurement unit configured to measure the output signal from each
output connection; a processing unit configured to process the
output signals so as to identify the micro-pixel generating a
signal or a center of mass of micro-pixels generating a signal, and
a global signal receiving unit comprising a preamplifier
electrically coupled in series with a capacitor, wherein the
capacitor is electrically coupled to the micro-pixels.
2. The device of claim 1, wherein the signal division network
further includes a quenching circuit comprising at least one of a
resistive network, a capacitive network, and an inductive network,
the quenching circuit being configured to terminate Geiger
discharges from the micro-pixels.
3. The device of claim 1, wherein the signal division network
further includes a two-dimensional array of elements.
4. The device of claim 1, wherein the signal division network
further includes two linear arrays of elements.
5. The device of claim 1, wherein the global signal receiving unit
is configured to obtain a global signal from a common
electrode.
6. The device of claim 5, wherein the global signal is obtained by
summing four position information signals.
7-10. (canceled)
11. The device of claim 1, wherein the signal output measurement
unit includes a resistive element and a preamplifier electrically
coupled to each of the four output connections.
12. A radiation detection system comprising: an integrated silicon
solid state photomultiplier (SSPM) device configured for operation
in a Geiger mode, the SSPM device including: a pixel unit including
an array of more than 2.times.2 p-n Geiger photodiodes (GPDs) on a
common substrate, a signal division network electrically coupled to
each GPD, the signal division network including four output
connections each for providing an output signal, means for
measuring the output signal from each output connection; means for
processing the output signals so as to identify the GPD generating
a signal or a center of mass of GPDs generating a signal, and means
for obtaining a global signal; and a scintillator optically coupled
to the SSPM device.
13. The system of claim 12 wherein the scintillator comprises
individual elements matched to the geometry of the integrated
silicon SSPM.
14. The system of claim 12 wherein the scintillator comprises a
pixilated microcolumnar scintillator.
15. The system of claim 12 wherein the scintillator comprises a
substantially continuous scintillator.
16. A radiation detection system comprising: an array of integrated
silicon SSPM devices operable in a Geiger mode, wherein each SSPM
device comprises a plurality of pixel units, each pixel unit
comprising an array of more than 2.times.2 p-n GPDs on a common
substrate, wherein the array of integrated SSPM devices comprises
an assembly of single pixels.
17. A combined PET/MRI imaging system, comprising: a main magnet
for use in MRI imaging; an RF coil disposed in the main magnet for
use in MRI imaging; and a PET scanner disposed between the main
magnet and the RF coil for use in PET imaging, the PET scanner
comprising the integrated silicon solid state photomultiplier
device of claim 1.
18. A method for detecting radiation, the method comprising:
providing a system including a pixel unit having an array of more
than 2.times.2 p-n micro-pixels on a common substrate, wherein the
system is operably coupled to a scintillator; providing an output
signal to each of four output connections, wherein each output
signal is generated based on signals generated from the
micro-pixels; measuring the output signal from each output
connection; processing the output signals so as to identify the
micro-pixel generating a signal or a center of mass of micro-pixels
generating a signal; obtaining a global signal from a common
electrode by summing four position information signals, wherein the
position information signals are based on the output signals; and
outputting the global signal to reconstruct an image for
display.
19. The method of claim 18, further comprising terminating Geiger
discharges from the micro-pixels.
Description
CROSS-REFERENCE TO RELATED APPLICATION
[0001] This application is a continuation of U.S. application Ser.
No. 12/580,172, filed Oct. 15, 2009, which claims the benefit of
priority under 35 U.S.C. .sctn.119(e) of U.S. Provisional
Application No. 61/185,169, filed Jun. 8, 2009, the entire contents
of which are incorporated herein by reference.
BACKGROUND OF THE INVENTION
[0003] The present invention relates to position sensitive
solid-state photomultipliers, and related systems and methods. More
specifically, the present invention relates to integrated silicon
solid-state photomultipliers and complementary metal-oxide
semiconductor (CMOS) avalanche photodiodes (APDs) operable in a
Geiger mode.
[0004] Photomultiplier tubes (PMTs) have been playing an important
role as photodetectors for the last several decades. PMTs, however,
are relatively expensive, bulky, have relatively low quantum
efficiency, come in specific fixed sizes, and are incompatible with
magnetic fields, the latter issue being relevant to the growing
interest in combined positron emission tomography (PET)/magnetic
resonance imaging (MRI) systems. PIN photodiodes and APDs, which
are compact, solid-state detectors, have previously been used to
build PET detectors. However, they have demanding operating
requirements, low gain, and poor timing capabilities.
[0005] Silicon solid state photomultipliers (SSPMs) are a promising
photodetection technology of increased recent interest. SSPM
technology is of interest for possible implementation in a wide
variety of applications, including PET, astroparticle physics and
gamma-ray astrophysics, high energy collider experiments, and dark
matter detection experiments. However, existing SSPM technologies
suffer from relatively high cost and complicated implementations,
as well as other deficiencies.
[0006] Accordingly, a need exists for SSPMs, and related devices,
systems, and methods, as well as SSPMs for use in various radiation
detection applications, including medical imaging applications.
BRIEF SUMMARY OF THE INVENTION
[0007] The present invention provides integrated silicon solid
state photomultiplier (SSPM) devices as well as related radiation
detection devices and imaging systems.
[0008] In one embodiment, an integrated silicon SSPM device is
operable in a Geiger mode. The SSPM device comprises a pixel unit
including an array of more than 2.times.2 p-n micro-pixels on a
planar substrate; a signal division network electrically coupled to
each micro-pixel, the signal division network including four output
connections each for providing an output signal; a signal output
measurement unit configured to measure the output signal from each
output connection; a processing unit configured to process the
output signals so as to identify the micro-pixel generating a
signal or a center of mass of micro-pixels generating a signal; and
a global signal receiving unit comprising a preamplifier
electrically coupled in series with a capacitor, wherein the
capacitor is electrically coupled to the micro-pixels.
[0009] In another embodiment, a radiation detection system is
provided. The radiation detection system comprises an integrated
silicon SSPM device configured for operation in a Geiger mode, and
a scintillator optically coupled to the SSPM device. The SSPM
device includes a pixel unit including an array of more than
2.times.2 p-n Geiger photodiodes (GPDs) on a common substrate; a
signal division network electrically coupled to each GPD, the
signal division network including four output connections each for
providing an output signal; means for measuring the output signal
from each output connection; means for processing the output
signals so as to identify the photodiode generating a signal or a
center of mass of photodiodes generating a signal; and means for
obtaining a global signal
[0010] In yet another embodiment, a radiation detection system
comprises an array of integrated silicon SSPM devices operable in a
Geiger mode, where each SSPM device comprises a plurality of pixel
units, and each pixel unit comprises an array of more than
2.times.2 p-n GPDs on a common substrate. Here, the array of
integrated SSPM devices comprises an assembly of single pixels.
[0011] In a further embodiment, a method for detecting radiation is
provided. The method includes: providing a system including a pixel
unit having an array of more than 2.times.2 p-n micro-pixels on a
common substrate, wherein the system is operably coupled to a
scintillator; providing an output signal to each of four output
connections, wherein each output signal is generated based on
signals generated from the micro-pixels; measuring the output
signal from each output connection; processing the output signals
so as to identify the micro-pixel generating a signal or a center
of mass of micro-pixels generating a signal; obtaining a global
signal from a common electrode by summing four position information
signals, wherein the position information signals are based on the
output signals; and outputting the global signal to reconstruct an
image for display.
[0012] For a fuller understanding of the nature and advantages of
the present invention, reference should be made to the ensuing
detailed description taken in conjunction with the accompanying
drawings. The drawings represent embodiments of the present
invention by way of illustration. The invention is capable of
modification in various respects without departing from the
invention. Accordingly, the drawings/figures and description of
these embodiments are illustrative in nature, and not
restrictive.
BRIEF DESCRIPTION OF THE DRAWINGS
[0013] FIG. 1 is a diagram of a radiation detection assembly.
[0014] FIG. 2 is a diagram of a photodetector/imaging device and a
data analysis/computer control system.
[0015] FIG. 3 illustrates a photodetector according to an
embodiment.
[0016] FIG. 4A illustrates a photodetector and portions of a signal
processing device according to an embodiment.
[0017] FIG. 4B illustrates a Geiger photodiode (GPD) according to
an embodiment.
[0018] FIG. 5A illustrates a photodetector and portions of a signal
processing device according to another embodiment.
[0019] FIG. 5B illustrates a GPD according to another
embodiment.
[0020] FIG. 6A illustrates a photodetector and portions of a signal
processing device according to a further embodiment.
[0021] FIG. 6B illustrates a GPD according to a further
embodiment.
[0022] FIG. 7 illustrates a diagram of a single photon emission
computed tomography (SPECT) imaging system.
[0023] FIG. 8A illustrates a diagram of a PET imaging system.
[0024] FIG. 8B illustrates an array of detectors for a PET imaging
system.
[0025] FIG. 9A illustrates a diagram of a combined PET/MRI imaging
system.
[0026] FIG. 9B illustrates a cross section of a combined PET/MRI
imaging system.
[0027] FIG. 10A shows a picture of a packaged a position sensitive
SSPM (PS-SSPM) chip.
[0028] FIG. 10B shows an enlarged view of a four-quadrant PS-SSPM
chip.
[0029] FIG. 11 shows the energy resolution at a varying bias
voltage for Q1 and Q4 PS-SSPMs.
[0030] FIG. 12A shows the timing distribution and resolution for a
Q1 PS-SSPM.
[0031] FIG. 12B shows the timing distribution and resolution for a
Q4 PS-SSPM.
[0032] FIG. 13A shows the spatial resolution for a Q1 PS-SSPM.
[0033] FIG. 13B shows the spatial resolution for a Q4 PS-SSPM.
[0034] FIG. 14 shows a picture of a LYSO scintillator array.
[0035] FIG. 15A shows an image produced by a Q1 PS-SSPM.
[0036] FIG. 15B shows an image produced by a Q4 PS-SSPM.
[0037] FIG. 16A shows a picture of a four-quadrant SSPM chip, where
each quadrant has an SSPM with unique design parameters.
[0038] FIG. 16B shows a picture of a four-quadrant SSPM chip, where
all quadrants have the same design parameters.
[0039] FIG. 17 shows a picture of an SSPM mounted on a
substrate.
[0040] FIG. 18 shows a picture of an MRI system with an inserted
lutetium yttrium oxyorthosilicate (LYSO) SSPM PET detector.
[0041] FIG. 19A shows a picture of a single event pulse output from
a PET detector arranged inside a magnetic field of an MRI system
with no MRI sequences.
[0042] FIG. 19B shows a picture of a single event pulse output from
a PET detector arranged inside a magnetic field of an MRI system
with a standard spin echo running.
[0043] FIG. 19C shows a picture of a single event pulse output from
a PET detector arranged inside a magnetic field of an MRI system
with a standard gradient echo sequence running.
[0044] FIG. 20 shows energy spectra from a PET detector arranged
inside a magnetic field of an MRI system.
[0045] FIG. 21A shows MRI images acquired without a PET
detector.
[0046] FIG. 21B shows MRI images acquired with an inactive PET
detector located in the MRI system.
[0047] FIG. 21C shows MRI images acquired with an active PET
detector located in the MRI system.
DETAILED DESCRIPTION OF THE INVENTION
[0048] The present invention includes methods, devices, and systems
including an integrated silicon solid state photomultiplier (SSPM)
device.
[0049] Some terminology used in this application is defined
below.
[0050] A single solid state device operable in Geiger mode is
referred to as a micro-pixel or a Geiger photodiode (GPD).
[0051] An integrated device that has more than 2.times.2 GPDs is
referred to as a solid state photomultiplier (SSPM) if there is no
resistive network for imaging. Such SSPMs can provide energy and
timing information but not imaging information. In some
embodiments, SSPMs may have 1000.times.1000 GPDs with a dimension
of 1.times.1 mm.sup.2 to a few cm.sup.2. A combined SSPM element is
referred to as a SSPM pixel.
[0052] With a resistive network for imaging, the device is referred
to as a position sensitive solid state photomultiplier (PS-SSPM),
which can provide energy, timing and position information.
[0053] A monolithic or assembled array of SSPMs is called a SSPM
(or PS-SSPM) array.
[0054] As described herein, the present invention includes SSPMs
comprising an array of micro-pixel p-n photodiodies. In some
embodiments, a size of the SSPM ranges from 20-50 .mu.m. The
photodiodes may be operated in Geiger mode using a relatively low
applied bias (<60 V). In Geiger mode, a diode is operated above
its breakdown threshold voltage, where a single electron-hole pair
can trigger a strong avalanche. In the case of such an event,
electronics may operate to reduce the voltage at the diode below
the threshold voltage for a short time, so that the avalanche is
stopped. The photodiode will then be operable once again to detect
further photons. Photodiodes operated in Geiger mode advantageously
increase a gain of a diode. For example, Geiger photodiodes may
provide a gain in a range of 10.sup.5 to 10.sup.6.
[0055] Each micro-pixel in the array may operate independently,
producing the same high gain output for a given micro-pixel
capacitance and bias, regardless of the number of incident photons
due to the Geiger mode operation. The micro-pixels in the array may
share a planar substrate. Accordingly, the outputs from all of the
micro-pixels, which may number as much as, e.g., 10.sup.3/mm.sup.2,
are summed together to produce a single macro SSPM analog output.
Energy information may be available where although each micro-pixel
provides a fixed signal independent of the number of photons
reaching it during its on-time, if the micro-pixel density is high
enough, then on average, each micro-pixel is triggered by a single
photon, and the number of micro-pixels producing a signal becomes
an indication of the total number of photons in a light pulse.
[0056] As further described herein, SSPM technology has been
developed using a complementary metal-oxide-semiconductor (CMOS)
process. This approach is attractive since it allows high precision
SSPM mass production at a low cost, and additionally offers the
possibility to integrate signal processing electronics with the
SSPM. As further described herein, discrete CMOS SSPMs, as well as
4.times.4 SSPM arrays, have been developed. Configured as an array,
the SSPM outputs may all be connected to resistive networks
providing position sensitive information on the micro-pixel level
while requiring only a limited number (e.g., four) of readout
channels to generate an image. These devices may not only indicate
where light is incident upon the SSPM sensing area, but also
provide spectroscopic and timing information as needed.
[0057] The SSPM technology described herein may be incorporated
into a variety of systems and devices. For example, the SSPM
technology described herein may be used in the detection and
spectroscopy of energetic photons (e.g., X-rays, gamma-rays, etc.).
Such detectors are commonly used, for example, in nuclear and
particle physics research, medical imaging, diffraction, non
destructive testing, nuclear treaty verification and safeguards,
nuclear non-proliferation monitoring, and geological exploration.
The SSPM technology may also be used in, for example, single photon
emission computed tomography (SPECT) imaging systems, x-ray
computed tomography scanner systems, positron emission tomography
(PET) systems, and combined PET/MRI systems.
[0058] FIG. 1 is a schematic diagram of a detector assembly 10. The
detector assembly 10 includes a scintillator 12 optically coupled
to a light photodetector/imaging device 14. The detector assembly
10 may also include a data analysis/computer control system 16 to
process information from the scintillator 12 and light
photodetector 14. The computer control system 16 may be separate
from the light photodetector 14 as illustrated in FIG. 1.
Alternatively, the computer control system 16 and the light
photodetector 14 may comprise a single unit. In use, the detector
assembly 10 detects energetic radiation emitted from a source 18.
The detector assembly 10 can be included, in whole or in part, in
detector and imaging systems, e.g., as described further below.
[0059] The data analysis/computer control system 16 can include,
for example, a module or system to process information (e.g.,
radiation detection information) from the photodetector 14. The
module or system can include, for example, a wide variety of
proprietary or commercially available computers, electronics, or
systems having one or more processing structures, a personal
computer, mainframe, or the like, with such systems often
comprising data processing hardware and/or software configured to
implement any one (or combination of) the method steps described
herein. Any software will typically comprise machine readable code
of programming instructions embodied in a tangible media such as a
memory, a digital or optical recording media, optical, electrical,
or wireless telemetry signals, or the like, and one or more of
these structures may also be used to transmit data and information
between components of the system in any of a wide variety of
distributed or centralized signal processing architectures.
[0060] The scintillator 12 may comprise structures and compositions
as described in, e.g., U.S. Pat. No. 7,129,494, U.S. Pat. No.
7,405,404, U.S. Pat. No. 7,365,333, U.S. Pat. No. 7,405,406, U.S.
Pat. No. 7,375,341, U.S. Pat. No. 7,361,901, U.S. patent
application Ser. No. 11/535,797, U.S. patent application Ser. No.
11/754,208, U.S. patent application Ser. No. 11/843,881, U.S.
patent application Ser. No. 11/773,356, U.S. patent application
Ser. No. 11/938,172, U.S. patent application Ser. No. 11/894,484,
U.S. patent application Ser. No. 12/334,351, U.S. patent
application Ser. No. 11/938,176, U.S. patent application Ser. No.
12/405,168, U.S. patent application Ser. No. 12/490,955, U.S.
patent application Ser. No. 12/497,436, U.S. Patent Application No.
61/230,970, all of which are incorporated by reference herein in
their entirety. For example, the scintillator 12 may comprise a
doped strontium iodide, where the dopant comprises europium,
cerium, or thallium. The scintillator 12 may comprise a
crystalline, ceramic, or polycrystalline ceramic form. The
scintillator 12 may comprise a plurality of elements such as
scintillation crystals that luminesce when excited by ionizing
radiation. The scintillator 12 may be a pixilated microcolumnar
scintillator. The scintillator 12 may be substantially
continuous.
[0061] The detector assembly 10, which can include the scintillator
12 and the photodetector 14, can be connected to a variety of tools
and devices, as mentioned previously. Various technologies for
operably coupling or integrating a radiation detector assembly
containing a scintillator to a detection device can be utilized in
the present invention, including various known techniques. The
detectors may also be connected to a visualization interface,
imaging equipment, or digital imaging equipment (e.g., pixilated
flat panel devices).
[0062] The detector assembly 10 may include means for permitting
radiation-induced scintillation light to pass out of the
scintillator 12 for measurement by the photodetector/imaging device
14. For example, the detector assembly 10 may include an optical
window at an end of a casing enclosing the scintillator 12, where
the end of the casing faces the photodetector/imaging device 14.
The window thus permits radiation-induced scintillation light to
pass out of the scintillator 12 for measurement by the
photodetector/imaging device 14, which is coupled to the
scintillator 12. The photodetector/imaging device 14 converts the
light photons emitted from the scintillator 12 into electrical
pulses that may be shaped and digitized by, for example, the
associated electronics. By this general process, radiation such as
gamma-rays can be detected.
[0063] FIG. 2 illustrates a photodetector/imaging device 14 and a
data analysis/computer control system 16 according to an
embodiment. The photodetector 14 includes a pixel unit 20
comprising a plurality of micro-pixels or GPDs. Each pixel may
comprise a light detecting element such as a GPD. Each GPD may
range in size from 20 to 40 .mu.m. The plurality of GPDs may be
arranged in a single array or into a plurality of arrays, and may
be provided at a density of around 10.sup.3/mm.sup.2. For example,
2.times.2 arrays, 4.times.4 arrays, 6.times.6 arrays, and the like
may be provided. Arrays having a size greater than 6.times.6 may
also be provided. The photodiodes may all share a common substrate,
and may be p-n photodiodes formed using a 0.8 .mu.m CMOS process.
Each photodiode may be operated in Geiger mode using a relatively
low applied bias (e.g., <60 V), and may operate independently to
produce a high gain (e.g., 10.sup.5 to 10.sup.6) output.
[0064] The photodetector/imaging device 14 need not be limited to
p-n photodiodes formed using a 0.8 .mu.m CMOS process, in a size
from 20 to 50 .mu.m, and provided at a density of around
10.sup.3/mm.sup.2, and operated in Geiger mode using a relatively
low applied bias. Rather, other types of photodiodes may be used.
For example, PIN photodiodes may be used. Other types of
fabrication processes may be used as known in the art. The size of
the photodiodes may be less than 20 .mu.m and greater than 50
.mu.m. The photodiodes may be provided at a density less than or
greater than 10.sup.3/mm.sup.2. The photodiodes may be operated in
other modes, such as photovoltaic mode and photoconductive
mode.
[0065] The photodetector 14 also includes a signal division network
22, where an electrical connection is provided between the signal
division network 22 and each of the photodiodes in the pixel unit
20. The signal division network 22 is electrically coupled to the
pixel unit 20 such that electric currents provided by each
photodiode are distributed over the signal division network 22. The
signal division network 22 may use passive quenching to terminate
Geiger discharges from the photodiodes of the pixel unit 20, where
the passive quenching may be facilitated via resistive elements
provided in the signal division network 22. As a result, the signal
division network may be a resistive network. In other embodiments,
the signal division network may be a capacitive network and/or an
inductive network. In some embodiments, active quenching elements
may also be used.
[0066] The signal division network 22 provides a plurality of
signals, such as signals A, B, C, and D, for measuring electric
currents provided by the pixel unit 20. The signal division network
22 includes output connections for each of the plurality of
signals, where each output connection provides one of the plurality
of signals. For example, four output connections may be provided,
where each output connection outputs one of the signals A, B, C,
and D. The signals A, B, C, and D may have current amplitudes that
are proportional to a location of an activated/firing photodiode or
proportional to a center of mass of activated/firing photodiodes.
The signal division network 22 may also provide a global signal for
communicating energy and timing information such as energy and
timing resolution.
[0067] The data analysis/computer control system 16 includes a
signal processing device 24 for receiving the plurality of signals
from the signal division network 22, such as signals A, B, C, D,
and Global, and for performing signal processing on the received
signals. The signal processing device 24 includes a global signal
receiving unit 26, a signal output measurement unit 28, and a
processing unit 30. The global signal receiving unit 26 receives
the global signal from a common electrode in the signal division
network 22. The global signal receiving unit 26 may include an AC
coupled charge sensing preamplifier for receiving the global
signal. The global signal is approximately equal to the sum of the
other signals, such as A, B, C, and D, provided by the signal
division network 22.
[0068] The signal processing device 24 also includes a signal
output measurement unit 28 for measuring the output signal from
each output connection of the signal division network 22. For
example, the signal output measurement unit 28 may measure the
output signals A, B, C, and D. The signal output measurement unit
28 may include charge sensitive preamplifiers, transimpedance
amplifiers and/or operational amplifiers for shaping and/or
applying a gain to the output signals A, B, C, and D so as to
produce conditioned output signals A', B', C', and D' from the
output signals A, B, C, and D, respectively. The A, B, C, and D,
and A', B', C', and D' refer to the position-sensitive outputs of
different embodiments of each of the photodiodes in the pixel unit
20.
[0069] The signal processing device 24 further includes a
processing unit 30 for processing output signals from the signal
division network 22 so as to identify the photodiode generating a
signal or so as to identify a center of mass of photodiodes
generating a signal. In an embodiment, the output signals A, B, C,
and D are processed. In another embodiment, the conditioned output
signals A', B', C' and D' are processed. As previously mentioned,
electric currents provided by each photodiode are distributed over
the signal division network 22 and have amplitudes that are
proportional to a location of an activated/firing photodiode or
proportional to a center of mass of activated/firing photodiodes.
Accordingly, the processing unit 30 analyzes either the output
signals A, B, C, and D or the conditioned output signals A', B',
C', and D' to calculate the location, such as an X and Y position,
of the activated/firing photodiode or the location of a center of
mass of photodiodes. For example, the processing unit 30 may
include Anger logic that uses the peak amplitude from the received
signals to provide spatial information on an event-by-event basis.
The processing unit 30 may then generate an image or cause an image
to be generated based on the calculated locations. As a result,
images can be generated using a scintillator 12 optically coupled
to a photodetector/imaging device 14 and a data analysis/computer
control system 16.
[0070] FIG. 3 illustrates a photodetector 14 according to an
embodiment. The photodetector 14 includes a pixel unit comprising a
plurality of micro-pixels 32. The photodetector 14 also includes a
signal division network comprising a plurality of resistive
components 34. According to the embodiment illustrated in FIG. 3,
sixteen micro-pixels 32 are provided. However, more or less
micro-pixels 32 may be provided. Each micro-pixel 32 may comprise a
photodiode operable in a Geiger mode. A resistive component 34 is
provided between every pair of micro-pixels 32. For example, a
resistive element 34 may be provided in series between each pair of
micro-pixels 32 that are arranged in a row. A resistive element 34
may also be provided in series between each pair of micro-pixels 32
that are arranged in a column. As a result, the signal division
network comprises a two-dimensional array of elements. The signal
division network may also comprise two or more linear arrays of
elements. In other embodiments, the signal division network may
comprise non-linear arrays.
[0071] The signal division network includes four output
connections, A, B, C, and D, where one output connection is
provided at each corner of the signal division network. The output
connections thus provide signals having current amplitudes that are
proportional to a location of an activated/firing photodiode or
proportional to a center of mass of activated/firing photodiodes.
The signals from the output connections may then be used by the
signal processing device 24 to determine the location of an
activated/firing photodiode or the location of the center of mass
of activated/firing photodiodes.
[0072] FIG. 4A illustrates a photodetector 14 and portions of a
signal processing device 24 according to an embodiment. The
photodetector 14 includes a pixel unit comprising a plurality of
micro-pixels 32. The photodetector 14 also includes a signal
division network comprising a plurality of resistive elements 34.
According to the embodiment illustrated in FIG. 4A, the
micro-pixels 32 have two terminals. The micro-pixels 32 arranged in
a row are coupled in parallel with one another and are each coupled
in parallel with a resistive element 34. A resistive element 34 is
also coupled in series between each pair of micro-pixels 32 that
are arranged in a row. The resistive elements 34 may each comprise,
for example, a 2 k.OMEGA. resistor. A bias voltage is applied to
one of the terminals of each micro-pixel 32. Although the
embodiment illustrated in FIG. 4A provides a 3.times.3 array of
micro-pixels 32 and corresponding signal division network, a
smaller or larger array may be provided. For example, a 2.times.2
array may be provided, as could a 4.times.4 array, 5.times.5 array,
and the like. According to an embodiment, arrays greater than
2.times.2 are provided.
[0073] As illustrated in FIG. 4A, a global signal receiving unit 26
is electrically coupled to the bias voltage and includes a
capacitor 36 and a preamplifier 38 arranged in series with one
another. The preamplifier 38 shapes and/or applies a gain to the
output of the capacitor 36 so as to provide the global signal,
which can subsequently be output to the processing unit 30. The
global signal receiving unit 26 is electrically coupled to the same
terminal of each of the micro-pixels 32. A signal output
measurement unit 28 is provided for measuring the output signal
from each output connection of the signal division network. The
signal output measurement unit 28 includes a pull-down resistor 40
and a preamplifier 42 for each output connection, where the
preamplifier 42 shapes and/or applies a gain to the output signal
at the output connection. For example, as illustrated in FIG. 4A,
the signal output measurement unit 28 shapes and/or applies a gain
to the output signal C so as to produce conditioned output signal
C'. Although the illustration in FIG. 4A shows the signal output
measurement unit 28 only corresponding to the output signal C', the
signal output measurement unit 28 may also correspond to the output
signals A', B', and D'. In some embodiments, the pull-down resistor
40 may be replaced with a pull-up resistor when the device is
biased with a negative supply.
[0074] FIG. 4B illustrates a micro-pixel 32 according to an
embodiment. The micro-pixel 32 includes a resistive element 44
arranged in series with a photodiode 46. The photodiode 46 is
operable in a Geiger mode, and the resistive element 44 passively
quenches, and distributes into the signal division network, the
Geiger discharge from the photodiode 46. The resistive element 44
may comprise, for example, a 200 k.OMEGA. resistor.
[0075] According to the embodiments illustrated in FIG. 4A and FIG.
4B, the processing unit 30 of the signal processing device 24 may
receive conditioned output signals from each of the output
connections of the signal division network. For example, the
processing unit 30 may receive signals A', B', C', and D'. The
processing unit 30 may then calculate the X and Y position for an
activated/firing photodiode in the array of micro-pixels 32 as:
X = ( A ' + B ' ) - ( C ' + D ' ) A ' + B ' + C ' + D ' Y = ( A ' +
D ' ) - ( B ' + C ' ) A ' + B ' + C ' + D ' ( 1 ) ##EQU00001##
[0076] FIG. 5A illustrates a photodetector 14 and portions of a
signal processing device 24 according to an embodiment. The labeled
elements illustrated in FIG. 5A which correspond to those
illustrated in FIG. 4A are similar, and thus further description is
omitted. However, in contrast with the embodiment illustrated in
FIG. 4A, the embodiment illustrated in FIG. 5A includes a pixel
unit comprising a plurality of micro-pixels 32 each having four
terminals. Further, the micro-pixels 32 arranged in a row are
coupled in parallel with one another, and the micro-pixels 32
arranged in a column are also coupled in parallel with one another.
A bias voltage is applied to a first one of the terminals of each
micro-pixel 32. A drain voltage is applied to a second one of the
terminals of each micro-pixel 32.
[0077] The row-column readout embodiment, illustrated in FIG. 5A,
maintains the orthogonality of the X and Y position signals. The
signal charge is equally divided and sent to two separate
charge-division networks: a charge-division network for the column
location, or X coordinate, and a charge-division network for the
row location, or Y coordinate. Resistive elements 34 are provided,
for each row of micro-pixels 32, in series between the two row
division network output connections. Similarly, resistive elements
34 are provided, for each column of micro-pixels 32, in series
between the two column division network output connections.
[0078] FIG. 5B illustrates a micro-pixel 32 according to an
embodiment. The labeled elements illustrated in FIG. 5B which
correspond to those illustrated in FIG. 4B are similar, and thus
further description is omitted. However, in contrast with the
embodiment illustrated in FIG. 4B, the embodiment illustrated in
FIG. 5B includes a pair of field effect transistors (FETs) 48
having gates G electrically coupled between the row or column
signal division network and the photodiode 46 of the micro-pixel
32. A drain voltage is applied to a drain D of each of the FETs 48.
The drain voltage may be, for example, equal to +3V. A source S of
one of the FETs 48 is electrically coupled to a row of the signal
division network. A source S of the other FET 48 is electrically
coupled to a column of the signal division network. The sources S
of the FETs 48 are electrically coupled to one another and between
the resistive element 44 and the photodiode 46 of the micro-pixel
32.
[0079] In this embodiment, the orthogonality of the X and Y
coordinates are maintained by FET transistors located at each
pixel. A charge pulse from a GPD element actives, equally, both FET
transistors. Each FET transistor is coupled to a linear resistor
network. One resistor network provides the column location, or X
coordinate, and the second resistor network provides the row
location, or Y coordinate. The role of the linear charge-division
network is similar to the embodiment illustrated in FIG. 6A. The
difference is that active components (e.g., two FET transistors)
are used to equally divide the signal charge into the two
charge-division networks, as opposed to the passive resistor
component used in FIG. 6B. With the use of active components, the
accuracy of the charge division does not depend on the relative
values of the quenching and network resistances.
[0080] According to the embodiments illustrated in FIG. 5A and FIG.
5B, the processing unit 30 of the signal processing device 24 may
receive conditioned output signals from each of the output
connections of the signal division network. For example, the
processing unit 30 may receive signals A', B', C', and D'. The
processing unit 30 may then calculate the X and Y position for an
activated/firing photodiode in the array of micro-pixels 32 as:
X = A ' - D ' A ' + D ' Y = B ' - C ' B ' + C ' ( 2 )
##EQU00002##
[0081] FIG. 6A illustrates a photodetector 14 and portions of a
signal processing device 24 according to an embodiment. The labeled
elements illustrated in FIG. 6A which correspond to those
illustrated in FIG. 5A are similar, and thus further description is
omitted. However, in contrast with the embodiment illustrated in
FIG. 5A, the embodiment illustrated in FIG. 6A includes a pixel
unit comprising a plurality of micro-pixels 32 each having three
terminals, where there is no drain voltage applied to a terminal of
the micro-pixels 32.
[0082] FIG. 6A shows an embodiment where passive components (e.g.,
resistive elements 44 in FIG. 6B) are used to equally divide the
charge into the two charge-division networks. The accuracy of
equally splitting the charge between the X and Y charge-division
networks improves when the collective resistance of the resistive
elements 44 is large relative to the total resistance of the charge
division networks, defined by the resistive elements 34. The
charge-division network is a linear array of resistive elements 44,
that divides the charge between two contacts for each
coordinate.
[0083] FIG. 6B illustrates a micro-pixel 32 according to an
embodiment. The labeled elements illustrated in FIG. 6B which
correspond to those illustrated in FIG. 4B are similar, and thus
further description is omitted. However, in contrast with the
embodiment illustrated in FIG. 4B, the embodiment illustrated in
FIG. 6B includes a pair of resistive elements 44. A terminal of one
of the resistive elements 44 is electrically coupled to a row of
the signal division network, while the other terminal of the
resistive element 44 is electrically coupled to a terminal of the
photodiode 46. A terminal of another one of the resistive elements
44 is electrically coupled to a column of the signal division
network, while the other terminal of the resistive element 44 is
electrically coupled to the terminal of the photodiode 46. A bias
voltage is applied to another terminal of the photodiode 46.
[0084] According to the embodiments illustrated in FIG. 6A and FIG.
6B, the processing unit 30 of the signal processing device 24 may
receive conditioned output signals from each of the output
connections of the signal division network. For example, the
processing unit 30 may receive signals A', B', C', and D'. The
processing unit 30 may then calculate the X and Y position for an
activated/firing photodiode in the array of micro-pixels 32 as
described above in equation (2).
[0085] FIG. 7 illustrates a basic configuration of a SPECT imaging
system 50. The system 50 can include configurations/components
commonly employed in known SPECT systems. As shown, the SPECT
imaging system 50 includes a patient or subject area 52 (positioned
subject shown for illustrative purposes), a detector assembly 54
and a computer control unit 56. The computer control unit 56 may
include circuitry and software for data acquisition, image
reconstruction and processing, data storage and retrieval, and
manipulation and/or control of various components/aspects of the
system. The detector assembly 54 can include a scintillator panel
or area including a scintillator material and a photodetector
assembly optically coupled to the scintillator material. The system
can include a single gamma-camera or detector in the detector
assembly or a plurality of detectors, with various configurations
and arrangements being possible. The detector assembly 54 and
subject area 52 may be movable with respect to each other, and may
include moving the detector assembly 54 with respect to the subject
area 52 and/or moving the subject area 52 with respect to the
detector assembly 54. In use, radiation detection includes
injecting or otherwise administering isotopes (including those
commonly employed in SPECT imaging) having a relatively short
half-life into the subject's body placeable within the subject area
52. The isotopes are taken up by the body and emit gamma-ray
photons that are detected by the detector assembly 54. SPECT
imaging is performed by using the detector assembly 54 to acquire
multiple images or projections (e.g., 2-D images), from multiple
angles. The computer control unit 56 is then used to apply image
reconstruction and processing, e.g., using a tomographic
reconstruction algorithm, to the multiple projections, yielding a
3-D dataset. This dataset may then be displayed as well as
manipulated to show different views, including slices along any
chosen axis of the body.
[0086] Light photodetector/imaging devices 14 including a pixel
unit 20 and a signal division network 22 of the present invention
can be utilized in the SPECT imaging system 50 and associated
imaging methods. For example, the light photodetector/imaging
devices 14 of the present invention can be incorporated into the
detector assembly 54. A signal processing device 24 of the present
invention can also be utilized in the SPECT imaging system 50 and
associated imaging methods. For example, the signal processing
devices 24 of the present invention can be incorporated into either
the detector assembly 54 or the computer control unit 56.
[0087] FIG. 8A illustrates a basic configuration of a PET imaging
system 60. In PET imaging, the PET imaging system 60 detects pairs
of gamma rays emitted indirectly by a positron-emitting
radionuclide (tracer), which is introduced into the subject's body.
Images of tracer concentration in 3-dimensional space within the
body are then reconstructed by computer analysis. PET imaging
systems and aspects of time of flight (TOF) PET imaging are further
described in commonly owned U.S. Pat. No. 7,504,634, which is
incorporated herein by reference in its entirety for all
purposes.
[0088] The PET imaging system 60 includes a PET camera system 62
having an array of radiation detectors 64, which may be arranged
(e.g., in polygonal or circular ring) around a patient area 66. In
some embodiments radiation detection begins by injecting or
otherwise administering isotopes with short half-lives into a
patient's body placeable within the patient area 66. As noted
above, the isotopes are taken up by target areas within the body,
the isotope emitting positrons that are detected when they generate
paired coincident gamma-rays. The annihilation gamma-rays move in
opposite directions, leave the body and strike the ring of
radiation detectors 64.
[0089] As shown in FIG. 8B, the array of detectors 64 includes an
inner ring of scintillators, including compositions as previously
described herein, and an attached ring of light detectors or
photomultiplier tubes. The scintillators respond to the incidence
of gamma rays by emitting a flash of light (scintillation) that is
then converted into electronic signals by a corresponding adjacent
photomultiplier tube or light detectors. A computer control unit or
system (not shown) records the location of each flash and then
plots the source of radiation within the patient's body. The data
arising from this process is usefully translated into a PET scan
image such as on a PET camera monitor by means known to those
having ordinary skill in the art.
[0090] Light photodetector/imaging devices 14 including a pixel
unit 20 and a signal division network 22 of the present invention
can be utilized in the PET imaging system 60 and associated imaging
methods. For example, the light photodetector/imaging devices 14 of
the present invention can be incorporated into the array of
radiation detectors 64. A signal processing device 24 of the
present invention can also be utilized in the PET imaging system 60
and associated imaging methods. For example, the signal processing
devices 24 of the present invention can be incorporated into either
the array of radiation detectors 64 or the computer control unit or
system (not shown).
[0091] In addition to gamma-ray imaging applications such as SPECT
and PET, many, indeed most, ionizing radiation applications will
benefit from the inventions disclosed herein. Specific mention is
made to X-ray fluoroscopy, X-ray cameras (such as for security
uses), and the like.
[0092] FIG. 9A illustrates a basic configuration of a combined
PET/MRI scanning system 80, and FIG. 9B illustrates a cross section
of the combined PET/MRI scanning system 80. The system 80 can
include configuration/components commonly employed in PET and MRI
systems and used simultaneously (e.g., simultaneous dynamic MRI,
simultaneous flow MRI, etc.) and further include a solid state
photomultiplier device of the present invention. As shown, the
PET/MRI scanning system 80 comprises an MRI scanner including a
main magnet 82 that has a hollow cylindrical geometry. The main
magnet 82 may embody the shape of a hollow cylinder. The main
magnet 82 generates a strong and uniform magnetic. Due to the
properties of the magnetic field, an MR image is typically acquired
within a central region of the main magnet 82. The MRI scanner also
includes a set of radio-frequency (RF) coils 84 typically located
within the central region. The RF coils 84 may transmit RF signals
to and receive RF signals from a subject during an MR imaging
process, where the subject is also typically located in the central
region.
[0093] The PET/MRI scanning system 80 also comprises a PET scanner
86 generally located within the hollow cylindrical shell of the
main magnet 82. The PET scanner 86 generally has a hollow
cylindrical geometry within which the RF coils 84 and subject are
located. The PET scanner 86 includes one or more detector
assemblies 88 and associated processing electronics, and a computer
control system (not shown) which may include circuitry and software
for image reconstruction, display, manipulation, post-acquisition
calculations, storage, data output, receipt, and retrieval. The
detector assemblies 88 may be provided in one or more rings around
the RF coils 84. Each ring may generate one slice of a PET image
for a subject. Hence, multiple rings may simultaneously generate
multiple slices of PET images for a subject. More specifically,
each ring of detector assemblies 88 collects high energy (e.g., 511
keV) annihilation photons produced by positron-electron
annihilations, wherein the positrons are emitted within the slice
of subject which is enclosed by the ring of scintillators. Next,
each of the high-energy photons that are collected by the
scintillators interacts with the scintillators to produce several
hundreds to thousands of low energy photons in the form of UV or
visible light photons which may subsequently be detected and
displayed as an image.
[0094] The detector assembly 88 can include a scintillator panel or
area including a scintillator material and a photodetector assembly
optically coupled to the scintillator material. The system further
includes (not shown) electronics coupled to the detector assembly
88 so as to output image data in response to radiation detection by
the scintillator.
[0095] Although not shown in FIG. 9A or FIG. 9B, the PET/MRI
scanning system 80 also includes a set of gradient coils which
generates field gradients onto the main magnetic field in the x, y,
and z directions. The field gradients are used to encode the
distance information in the space where the subject is located. In
one embodiment, the set of gradient coils is situated so that they
enclose the PET scanner 86. Generally, a PET/MRI scanning system 80
is constructed so that the PET scanner 86 is inserted in the open
space between the inner surface of main magnet 82 and RF coils
84.
[0096] Light photodetector/imaging devices 14 including a pixel
unit 20 and a signal division network 22 of the present invention
can be utilized in the combined PET/MRI scanning system 80 and
associated imaging methods. For example, the light
photodetector/imaging devices 14 of the present invention can be
incorporated into the detector assembly 88. A signal processing
device 24 of the present invention can also be utilized in the
combined PET/MRI scanning system 80 and associated imaging methods.
For example, the signal processing devices 24 of the present
invention can be incorporated into either the detector assembly 88
or the computer control system (not shown). Incorporating light
photodetector/imaging devices 14 and signal processing devices 24
according to the present invention in combined PET/MRI scanning
systems 80 is particularly advantageous since the solid state
photodiodes have a low sensitivity to magnetic fields and thus a
PET imaging system can be incorporated into an MRI scanning system
without detrimentally affecting an operation of the MRI scanning
system.
[0097] Systems and methods of the present invention as described
above are illustrative, and alternate configurations and
embodiments will be included. The present invention may include
modifications as well as combinations of imaging systems as
described, such as combined imaging systems--e.g., combined SPECT
MRI systems, and the like.
Examples
A. CMOS Position Sensitive SSPM
[0098] In one embodiment, "p on n" position sensitive solid state
photomultipliers (PS-SSPMs) using a high-voltage 0.8 .mu.m CMOS
process were developed, which were available through Metal Oxide
Semiconductor Implementation Service (MOSIS). MOSIS is a relatively
low cost production service for semiconductor development, which is
well suited for small volume prototyping and production runs. The
PS-SSPM had five output signals associated with it. There were four
output signals that provide X-Y spatial information when compared
amongst each other using Anger logic, and a PS-SSPM global signal,
which is approximately the sum of the four spatial signals. The
global signal can provide energy and timing information. Each
individual micro-pixel discharge output was connected to a
resistive network, where the total PS-SSPM current signal was
shared amongst four electrodes located at the network corners, or
end points. Depending on where the firing micro-pixels are,
relative to the four electrode locations in the resistive network,
the micro-pixel's current signal will be distributed over the
network and its amplitude, measured at the electrodes, will be
proportional to the firing micro-pixel location. By analyzing the
four signals using Anger logic, the X and Y position can be
calculated. So with this signal processing technique, it is
possible to generate an image using a finely segmented scintillator
array coupled to the PS-SSPM or some other incident focused light
source. According to the basic PS-SSPM concept, micro-pixels are
connected via a resistive network, and the signal charge is
collected from the network corners. Anger logic according to an
embodiment uses the peak amplitude from the four outputs to provide
spatial information on an event-by-event basis. In one embodiment,
four different PS-SSPM designs were provided, each with a different
resistor network scheme, along with different micro-pixel
geometries. Each PS-SSPM design used passive quenching to terminate
the micro-pixel Geiger discharge. Table 1 lists the micro-pixel
parameters for each PS-SSPM. One of the PS-SSPM designs, quadrant 2
(Q2), failed to produce any meaningful data due to issues believed
to be related to its network resistors. Despite their resistor
network differences, the PS-SSPMs all basically use the same
readout technique, as described above, to determine the event X-Y
location. The different resistor network designs for quadrant 1
(Q1), quadrant 3 (Q3), and quadrant 4 (Q4) are described in
forthcoming sections.
TABLE-US-00001 TABLE 1 Parameters for four PS-SSPMs. PS-SSPM
Quadrant Parameters Parameter Q1 Q2 Q3 Q4 Number of 1089 900 255
900 micro- pixels micro- 30 .times. 30 .mu.m.sup.2 30 .times. 30
.mu.m.sup.2 67 .mu.m .times. 38 .mu.m 30 .times. 30 .mu.m.sup.2
pixel dimen- sions micro- 44 .times. 44 .mu.m.sup.2 49 .times. 49
.mu.m.sup.2 .sup. 93 .times. 93 .mu.m.sup.2 47 .times. 47
.mu.m.sup.2 pixel pitch dimen- sions Fill 46% 37% 29% 41%
Factor
[0099] The four 1.5.times.1.5 mm.sup.2 PS-SSPM variations were all
fabricated on one 3.times.3 mm.sup.2 chip, where each PS-SSPM could
be independently operated and read out. The PS-SSPM chip was
packaged on a ceramic 145 pin grid array (PGA). FIG. 10A
illustrates the packaged PS-SSPM chip. FIG. 10B provides an
enlarged view of the PS-SSPMs provided in each of the four
quadrants.
[0100] A custom printed circuit board, with a mounted ZIF socket
for the PGA, was fabricated in order to easily operate and readout
the PS-SSPMs. The printed circuit board had a four-position switch
to select which PS-SSPM had its output signals routed into the
on-board charge sensitive preamplifiers.
Quadrant 1
[0101] The PS-SSPM fabricated at Q1 comprised a photodetector 14
and portions of a signal processing device 24 as illustrated in
FIG. 4A and FIG. 4B. Each micro-pixel in Q1 is passively quenched
with a 200 k.OMEGA. resistor and had its Geiger discharge injected
into a resistor grid network. This network connects each
micro-pixel output with its neighbors' output via 2 k.OMEGA.
resistors. Four electrodes, each located at the network corners,
collect the distributed charge so that charge sensing preamplifiers
can then produce the spatial output signals for the PS-SSPM. The X
and Y position for each event can be calculated using the shaped
preamplifier outputs and equation (1). There is also a charge
sensing preamplifier AC coupled to the PS-SSPM bias. This
preamplifier produces the global PS-SSPM signal that can be used
for energy and timing information.
Quadrant 3
[0102] The PS-SSPM fabricated at Q3 comprised a photodetector 14
and portions of a signal processing device 24 as illustrated in
FIG. 5A and FIG. 5B. Like Q1, each micro-pixel in this design also
had a 200 k.OMEGA. resistor to passively quench the Geiger
discharge. At the output of each micro-pixel, are two field effect
transistors (FETs). A+3 V drain voltage was applied to both FETs.
The respective FET output was fed into a single row or column of
resistors whose ends have charge sensing preamplifiers connected.
The location along the row and column where the Geiger discharge is
inserted provides the X and Y position information using equation
(2). Q3 also has a global signal that was taken from the bias
connection via an AC coupled charge sensing preamplifier.
Quadrant 4
[0103] The PS-SSPM fabricated at Q4 comprised a photodetector 14
and portions of a signal processing device 24 as illustrated in
FIG. 6A and FIG. 6B. The micro-pixel readout design in Q4 was much
like that of Q3, with the exception of the two FETs used at the
micro-pixel output. In this case, each micro-pixel simply has two
outputs for its Geiger discharge, each path possessing a 200
k.OMEGA. resistor for passive quenching. Again, each respective
micro-pixel output is fed into a single row and column of resistors
whose ends have connected charge sensing preamplifiers. Where along
the row of column the current Geiger discharge is inserted, will
provide the X and Y position information using equation (2). This
PS-SSPM also has a global signal.
Quadrant One and Quadrant Four PS-SSPM Characterizations
Energy Resolution Using LYSO
[0104] According to an embodiment, the energy resolution for Q1 and
Q4 was measured at 511 keV using a 1.times.1.times.20 mm.sup.3 LYSO
scintillator wrapped with several layers of Teflon tape. The
scintillator was coupled to each quadrant individually using
optical grease (Rexon, RX-688) and irradiated with .sup.22Na (511
keV). The global signal was used from each of the quadrants to
record the pulse height distribution. The global signal was fed
into a charge sensitive preamplifier and then shaped at 0.25
.mu.sec using a spectroscopy amplifier (Canberra, model 2020). An
MCA (Amptek, MCA-8000A) recorded and histogrammed the shaped
signals for analysis. The energy resolution (at 511 keV) for each
quadrant was measured in this way as the PS-SSPM bias was increased
in 1 V increments from 28 V to 35 V. Breakdown occurs at 27 V for
all quadrants. FIG. 11 shows the results. The energy resolution for
Q1 and Q4 were approximately equal from 27 V to 34 V. At 34 V the
FWHM energy resolution was 13.3% and 13.9% for Q1 and Q4
respectively.
Coincidence Timing Resolution Using LYSO
[0105] Using the same 1.times.1.times.20 mm.sup.3 LYSO
scintillator, the coincidence timing resolution was measured at 511
keV (.sup.22Na) for the Q1 and Q4 quadrants compared with a PMT
(Hamamatsu, H6533) coupled to a LYSO scintillator. A Canberra model
2003T preamplifier, which has a fast timing output, was AC coupled
to the PS-SSPM global signal. Standard NIM electronics were used to
perform the timing measurement. The LYSO-PMT provided the timing
analyzer start signal, while the LYSO-SSPM provided the stop
signal. Using this setup, the coincidence timing resolution was
measured for each of the quadrants. The FWHM timing resolution for
Q1 at 32 V was 2.1 nsec and Q4 at 32 V was 1.0 nsec. No energy
gating was used. FIG. 12A and FIG. 12B show the timing distribution
and resolution for Q1 and Q4, respectively.
Laser Based Spatial Resolution
[0106] According to an embodiment, the intrinsic spatial resolution
of Q1 and Q4 was measured using a pulsed 635 nm diode laser
(Thorlabs, HS9-635), driven by a pulse generator (Tektronix PG508),
focused into an approximately 15 .mu.m diameter beam spot that was
incident upon the middle of the quadrant's sensing area. The laser
intensity was approximately 1000 photons/pulse, with a pulse width
of 50 nsec at a frequency of approximately 1 kHz. For each laser
pulse, the four spatial output signals were shaped at 0.25 .mu.sec
by Canberra amplifiers (model 2020) and then sent to a sample and
hold circuit before being readout and digitized by our data
acquisition card (Keithley, DAS-1802HC). The X-Y position was
calculated for each event using equation (1) or (2) (depending on
which quadrant was being examined) and recorded in list-mode. Ten
thousand pulses were recorded by each quadrant and the calculated
X-Y positions were histogrammed. The histogrammed position
distribution FWHM defined the spatial resolution. FIG. 13A and FIG.
13B show the spatial resolution for Q1 and Q4, respectively. The
measured X-Y spatial resolution for Q1 and Q4 at 28 V bias were
both approximately 70 .mu.m.
Scintillator Array Imaging
[0107] According to an embodiment, imaging studies were performed
using a LYSO scintillator array. The LYSO array is illustrated in
FIG. 14 and was populated with 500.times.500 .mu.m.sup.2 pixels.
The array was coupled to the quadrants individually using optical
grease from Rexon. The array was larger than the quadrant sensing
area; however, for a simple imaging demonstration this was not a
concern. The detector and LYSO array were uniformly irradiated with
511 keV gamma-rays from .sup.22Na. The four PS-SSPM output signals
were processed in the same fashion as for the spatial resolution
measurements. Here, the data acquisition was triggered by the
PS-SSPM global signal. Given the 1.5.times.1.5 mm.sup.2 area for
each quadrant and the pixel dimensions, a 3.times.3 image is the
largest one can achieve. FIG. 15A and FIG. 15B show the array
images produced by Q1 and Q4, respectively. The LYSO array image
for Q1 shows a pincushion distortion, which is typical for similar
imaging sensors or systems that use Anger logic, such as equation
(1), to build an image. Note that Q4, which does not use a
resistive grid network, but rather uses a resistive row-column
readout, does not show a distortion.
B. LYSO-SSPM PET Detector for Combined PET/MRI Applications
[0108] According to an embodiment, a novel solid SSPM was designed
using standard CMOS technology and evaluated the SSPM for combined
PET/MRI systems.
CMOS SSPMs
[0109] According to an embodiment, high fill factor 2.times.2
arrays of SSPMs were built with each SSPM measuring 1.5.times.1.5
mm.sup.2. The design details for these SSPMs are provided in Table
2. FIG. 16A shows a chip that has four separate SSPMs, each with
1.5.times.1.5 mm.sup.2 area but with different design parameters.
In other words, each SSPM in a quadrant has unique design
parameters. These SSPMs have square micro-pixels (30-50 .mu.m
size). FIG. 16B shows a chip that also has four separate SSPMs,
each with 1.5.times.1.5 mm.sup.2 area but in this case each
quadrant has an identical design and as a result the whole chip can
be operated as a single 3.times.3 mm.sup.2 SSPM with .about.1750
pixels.
TABLE-US-00002 TABLE 2 Parameters for CMOS SSPMs. CMOS SSPM
Parameters Active micro-pixel Size # of Area (.mu.m)/micro-pixel
micro- Fill Operating (mm.sup.2) Pitch (.mu.m) pixels Factor Gain
Bias 1.5 .times. 1.5 30/56 576 29% .sup. 1 .times. 10.sup.6 30 V
(square) 1.5 .times. 1.5 30/43 961 49% .sup. 1 .times. 10.sup.6 30
V (square) 1.5 .times. 1.5 50/76 324 43% 3.5 .times. 10.sup.6 30 V
(square) 1.5 .times. 1.5 50/63 441 63% 3.5 .times. 10.sup.6 30 V
(square) 3 .times. 3 50/63 ~1750 63% 3.5 .times. 10.sup.6 30 V
(square)
Evaluation of LYSO-SSPM Detector for PET/MRI Studies
[0110] According to an embodiment, a PET detector was built,
consisting of a 1.5.times.1.5.times.10 mm.sup.3 LYSO crystal,
coupled with optical grease to a 1.5.times.1.5 mm.sup.2 SSPM with
49% fill-factor (described in the second row of Table 1). The SSPM
was mounted on a ceramic substrate and then placed on a 15.times.25
mm readout board. The readout board was made with non-magnetic
material and was finally mounted on a 45 cm long carbon fiber strip
as shown in FIG. 17. The detector was positioned in the annulus
between the RF and the gradient coils of UC Davis' 7T Bruker
Biospec MRI system, exactly where it would need to be for
simultaneous PET and MRI studies. FIG. 18 is a photograph of an MRI
system with an inserted LYSO-SSPM PET detector.
Effect of MRI on a LYSO-SSPM PET Detector
[0111] According to an embodiment, a .sup.68Ge gamma source was
placed in the cavity of the RF coil and at the center of the field
of view of an MRI system. The PET detector, as illustrated in FIG.
18, was biased at 30V. FIG. 19A, FIG. 19B, and FIG. 19C show single
event pulses from the output of the amplifier, digitized using a
fast oscilloscope when the detector was inside the magnetic field
(without any MRI sequences), when a standard spin echo was running,
and when a standard gradient echo sequence was running,
respectively. The PET signals were not distorted, nor was there a
visible baseline shift or increased noise.
[0112] FIG. 20 shows energy spectra recorded when the detector was
inside the magnet without any sequence and with spin echo and
gradient echo sequences on. A slight change (less than 2%) in the
location of the 511 keV photopeak was observed with and without MRI
sequences. The measured energy resolution was .about.12% (FWHM)
with and without MRI sequences, suggesting that there was no
significant loss in signal amplitude or increase in noise.
Effect of LYSO-SSPM PET Detector on MRI Images
[0113] According to an embodiment, a structured cylindrical MRI
phantom containing Magnevist.RTM. in water was imaged to assess the
effect of the PET insert on the MRI data acquisition. FIG. 21A
shows MRI images acquired without a PET detector. In the top image
a spin sequence was on, and in the bottom image a gradient echo
sequence was on. FIG. 21B shows MRI images acquired where a PET
detector was included in the MRI system and the PET detector was
turned off. In the top image a spin sequence was on, and in the
bottom image a gradient echo sequence was on. FIG. 21C shows MRI
images acquired where a PET detector was included in the MRI system
and the PET detector was powered on. In the top image a spin
sequence was on, and in the bottom image a gradient echo sequence
was on.
[0114] The images were visually inspected for the presence of
artifacts. No obvious artifacts can be observed in the resulting
phantom images when compared with the ones acquired without the PET
detector insert. This suggests that there was no significant
interference due to the PET detector on the magnetic field or RF
pulses of the MRI scanner. This investigation shows promise for
MR-compatible PET systems based on SSPM technology.
[0115] Although the invention has been described with reference to
the above examples, it will be understood that modifications and
variations are encompassed within the spirit and scope of the
invention. Accordingly, the invention is limited only by the
following claims along with their full scope of equivalents.
* * * * *