U.S. patent application number 10/706821 was filed with the patent office on 2006-11-23 for method and apparatus for improving pet detectors.
Invention is credited to Dario B. Crosetto.
Application Number | 20060261279 10/706821 |
Document ID | / |
Family ID | 37301210 |
Filed Date | 2006-11-23 |
United States Patent
Application |
20060261279 |
Kind Code |
A1 |
Crosetto; Dario B. |
November 23, 2006 |
METHOD AND APPARATUS FOR IMPROVING PET DETECTORS
Abstract
The present invention is directed to a system, method and
software program product for implementing an efficient,
low-radiation 3-D Complete-Body-Screening (3D-CBS) medical imaging
device which combines the benefits of the functional imaging
capability of PET with those of the anatomical imaging capability
of CT. The present invention enables a different detector assembly,
and together they enable execution of more complex algorithms
measuring more accurately the information obtained from the
collision of the photon with the detector. The present invention
overcomes input and coincidence bottlenecks inherent in the prior
art by implementing a massively parallel, layered architecture with
processor separate stacks for handling each channel. The prior art
coincidence bottleneck is overcome by limiting coincidence
comparisons to those with a time stamp occurring within a
predefined time window. The increased efficiency provides the
bandwidth necessary for increasing the throughput even more by
extending the FOV to over one meter in length and the execution of
even more complex algorithms.
Inventors: |
Crosetto; Dario B.; (DeSoto,
TX) |
Correspondence
Address: |
JONES DAY
P.O. BOX 660623
DALLAS
TX
75266-0623
US
|
Family ID: |
37301210 |
Appl. No.: |
10/706821 |
Filed: |
November 10, 2003 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
60424933 |
Nov 9, 2002 |
|
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Current U.S.
Class: |
250/367 |
Current CPC
Class: |
G01T 1/2985 20130101;
G01T 1/1642 20130101; G01T 1/1615 20130101; A61B 6/037 20130101;
G01T 1/1612 20130101 |
Class at
Publication: |
250/367 |
International
Class: |
G01T 1/202 20060101
G01T001/202 |
Claims
1. A device for detecting subatomic particles comprising: a
detector assembly, said detector assembly comprising: a first
plurality of transducers; and at least one detector crystal
optically coupled to at least some of said first plurality of
transducers, wherein a single detector crystal in said at least one
detector crystal has a plurality of slits, each slit being
approximately equivalent in length as each other slit and said
plurality of slits being oriented parallel to the optical axis of
said at least some of said first plurality of transducers; a
plurality of processors, said plurality of processors being
arranged in a plurality of substantially parallel layers, and some
processors in said plurality of processors being capable of
communicating with processors to each lateral side in one
substantially parallel layer, receiving communication from a
processor in a second substantially parallel layer and transmitting
to a third processor in a third substantially parallel layer; and
wherein each transducer in the first plurality of transducers is in
electrical communication with at least one processor in the
plurality of processors.
2. The device recited in claim 1 above, wherein said at least one
detector crystal forms a first side and an opposing second side and
wherein said first plurality of transducers are optically coupled
to said first side, the detector assembly further comprising: a
second plurality of transducers, said second plurality of
transducers optically coupled to the at least one detector crystal,
the second plurality of transducers being coaxial with at least
some of the first plurality of transducers, wherein said second
plurality of transducers are optically coupled to said second side,
and wherein a surface area of a face of a transducer in said second
plurality of transducers is smaller than a surface area of a face
of a transducer in said first plurality of transducers; and wherein
each transducer in the second plurality of transducers is in
electrical communication with at least one processor in said
plurality of processors.
3. The device recited in claim 2 above, wherein a light guide is
optically coupled between a detector crystal in the at least one
detector crystal and a transducer in the second plurality of
transducers.
4. The device recited in claim 2 above, wherein at least one of the
plurality of processors performs a depth of interaction
calculation.
5. The device recited in claim 1 above, the device further
comprising a pyramidal funneling structure, said pyramidal
funneling structure comprising a plurality of funnel input
processors, and wherein a funnel input processor in the plurality
of funnel input processors is coupled to a processor in the
plurality of processors of claim 1.
6. The device recited in claim 1 above, wherein a transducer in
said first plurality of transducers is a photomultiplier (PMT) or
an avalanche photodiode (APD).
7. The device recited in claim 2 above, wherein a transducer in
said second plurality of transducers is a photomultiplier (PMT), an
avalanche photodiode (APD), or a photodiode.
8. The device recited in claim 1 above, wherein a detector crystal
in the at least one detector crystal is a bismuth germinate (BGO)
crystal or a sodium iodate (NaI) crystal.
9. The device recited in claim 2 above, wherein said at least one
detector crystal defines a barrel around a patient and wherein the
first plurality of transducers are arranged on an exterior face of
said barrel and said second plurality of transducers are arranged
on an interior face of said barrel.
10. The device recited in claim 9 above, wherein said at least one
detector crystal is a single detector crystal.
11. The device recited in claim 9 above, wherein said barrel is
segmented into sectors, and wherein said at least one detector
crystal consists of four separate detector crystals, each detector
crystal in said four separate detector crystals occupying a
different sector of said barrel.
12. The device recited in claim 9 above, wherein said barrel is
segmented into sectors, and wherein said at least one detector
crystal consists of two separate detector crystals, each detector
crystal in said two separate detector crystals occupying a
different sector of said barrel.
13. The device recited in claim 1 above, wherein a processor in
said plurality of processors is a FPGA.
14. The device recited in claim 1 above, wherein a processor in
said plurality of processors is an ASIC.
15. The device recited in claim 1 above, wherein the timing of a
processor in said plurality of processors is provided by two
in-phase clocks at 20 MHz and 40 MHz.
16. The device recited in claim 2 above, wherein a first transducer
in the first plurality of transducers is in electrical
communication with a first processor in said plurality of
processors; and a second transducer in said second plurality of
transducers is in electrical communication with said first
processor.
17. The device recited in claim 16 above, wherein said first
processor is in a first substantially parallel layer in said
plurality of substantially parallel layers and wherein said first
processor is in electrical communication with four other processors
in said first substantially parallel layer.
18. The device recited in claim 17, wherein said first processor is
in electrical communication with a second processor in a second
substantially parallel layer in said plurality of substantially
parallel layers.
19. The device recited in claim 1, wherein the subatomic particles
are photon pairs.
Description
CROSS REFERENCES TO RELATED APPLICATIONS
[0001] The present application is related to and claims priority
from the co-pending U.S. provisional patent application entitled
"METHOD AND APPARATUS FOR IMPROVING PET DETECTORS" having
application No. 60/424,933 filed on Nov. 9, 2002 and is
incorporated by reference herein by reference in its entirety.
[0002] The present application is also related to the following
patent applications:
[0003] U.S. Pat. No. 5,937,202 filed Feb. 15, 1996 entitled
"High-Speed, Parallel, Processor Architecture for Front-End
Electronics, Based on a Single Type of ASIC, and Method Use
Thereof," (hereinafter U.S. patent '202).
[0004] U.S. patent application Ser. No. 09/506,207 filed Feb. 15,
2000 entitled "Method and Apparatus for Extending Processing Time
in One Pipeline Stage," (hereinafter U.S. application '207), which
claims priority from: U.S. Provisional Patent Application No.
60/120,194 filed Feb. 16, 1999; U.S. Provisional Patent Application
No. 60/112,130 filed Mar. 12, 1999; U.S. Provisional Patent
Application No. 60/129,393 filed Apr. 15, 1999; U.S. Provisional
Patent Application No. 60/132,294 filed May 3, 1999; U.S.
Provisional Patent Application No. 60/142,645 filed Jul. 6, 1999;
U.S. Provisional Patent Application No. 60/143,805 filed Jul. 14,
1999; U.S. Provisional Patent Application No. 60/154,153, Sep. 15,
1999; U.S. Provisional Patent Application No. 60/161,458 filed Oct.
25, 1999; U.S. Provisional Patent Application No. 60/164,694 filed
Nov. 10, 1999; and U.S. Provisional Patent Application No.
60/170,565 filed Dec. 14, 1999.
[0005] U.S. patent application Ser. No. 10/185,904 filed Jun. 27,
2002 entitled "Method and Apparatus for Whole-Body,
Three-Dimensional Dynamic PET/CT Examination," (hereinafter U.S.
application '904), claiming priority from U.S. Provisional Patent
Application No. 60/301,545 filed Jun. 27, 2001; and U.S.
Provisional Patent Application No. 60/309,018 filed Jul. 31,
2001.
[0006] U.S. patent application Ser. No. 10/296,532 filed Nov. 25,
2002 entitled "Method and Apparatus for Anatomical and Functional
Medical Imaging," (hereinafter U.S. application '532), which claims
priority from: PCT/US01/15671 filed May, 15, 2001; U.S. Provisional
Patent Application No. 60/204,900 filed May 16, 2000; U.S.
Provisional Patent Application No. 60/215,667 filed Jun. 30, 2000;
U.S. Provisional Patent Application No. 239,543 filed Oct. 10,
2000; U.S. Provisional Patent Application No. 60/250,615 filed Nov.
30, 2000; U.S. Provisional Patent Application No. 60/258,204 filed
Dec. 22, 2000; and U.S. Provisional Patent Application No.
60/261,387 filed Jan. 15, 2001.
[0007] U.S. patent application Ser. No. 10/376,024 filed Feb. 26,
2003 entitled "Method And Apparatus For Determining Depth of
Interactions in a Detector for Three-Dimensional Complete Body
Screening," (hereinafter U.S. application '024), claiming priority
from U.S. Provisional Patent Application No. 60/360,301 filed Feb.
26, 2002.
[0008] U.S. patent application Ser. No. 10/453,255 filed Jun. 2,
2003 entitled "Gantry for Geometrically Configurable and
Non-Configurable Positron Emission Tomography Detector Arrays,"
(hereinafter U.S. application '255), claiming priority from U.S.
Provisional Patent Application 60/385,140 filed Jun. 2, 2002.
[0009] The above-identified patent applications are incorporated by
reference herein by reference in their entireties.
BACKGROUND OF THE INVENTION
[0010] 1. Field of the Invention
[0011] The present invention relates to positron emission
tomography (PET). More particularly, the present invention relates
to PET detectors used therein.
[0012] 2. Description of Related Art
[0013] The use of positron emissions for medical imaging has been
well documented from the early 1950s, see "A History of Positron
Imagining," Brownell, Gordon, presented on Oct. 15, 1999,
Massachusetts General Hospital and available at
http://neurosurgery.mgh.harvard.edu/docs/PEThistory.pdf, which is
incorporated herein by reference in its entirety. PET imaging has
advantages over other types of imaging procedures. Generally, PET
scanning provides a procedure for imaging the chemical
functionality of bodily organs rather than imaging only their
physical structure, as is commonly available with other types of
imaging procedures such as X-ray, Computerized Tomography (CT), or
Magnetic resonance imaging (MRI). PET scanned images allow a
physician to examine the functionality of the heart, brain, and
other organs, as well as diagnosing disease groups which cause
changes in the cells of a body organ or in the manner in which they
grow, change, and/or multiply out of control, such as cancers.
[0014] Positron Emission Tomography (PET) is a medical imaging
technique that involves injecting a natural compound, such as sugar
or water, labeled with a radioactive isotope into a patient's body
to reveal internal biological processes. As the isotope (positron)
circulates within the patient's body. The positron annihilates with
and electron and emits pairs of photons in diametrically opposed
directions (back-to-back). A PET device is made of a set of
detectors coupled to thousands of sensors that surround the human
body. These detectors (crystals) capture the photons emitted by the
isotope from within the patient's body at a total rate of up to
hundreds of millions per second, while the sensors (transducers
such as PMTs) convert them to electrical signals, and send the
signals to the electronics.
[0015] Other applications for detecting particles (photons,
electrons, hadron, muon and jets) are well known, such as with
regard to experiments in high energy physics. While particle
detection in high energy physics and medical imaging have some
common ground, differences between the disciplines are sticking. On
distinction between the usages is that the detectors used in
medical imaging are approximately 200 times smaller than the larger
detectors employed in high-energy physics applications, and what is
more, medical imaging PET applications require the identification
of only a single type of particle, the photon.
[0016] Typically, prior art Positron Emission Tomography (PET)
devices require the injection into the patient's body of a
radiation dose that is 10 to 20 times the maximum radiation dose
recommended by the International Commission on Radiological
Protection (ICRP). This amount is necessary because, at best, prior
art PET devices detect only detect two photons out of 10,000
emitted in the patients' body. Currently, the largest manufacturers
of PET (General Electric Company and Siemens AG (ADR)) which
command in excess of 90% of the world market, are manufacturing two
different PET (PET/CT) systems with very similar performance and
are selling them at very similar prices. However, although the
price and performance of the systems from the different
manufacturers are comparable, one manufacturer's system (Siemens)
uses nearly ideal crystal detectors, while contrastingly, the other
manufacturer's system (General Electric) uses cheaper, lower
quality crystal detectors with slower decay time. Consequently, the
manufacturer using the cheaper, lower cost detectors, expends on
the order of only 10% the price of the ideal crystals used in their
competitor's systems. Thus, the question arises as to how it could
be that, even though one manufacturer uses crystal detectors that
are ten times more expensive that the other manufacturer, the price
and performance of the two PET systems from the different
manufacturers are very comparable.
[0017] Anecdotally, the present inventor has analyzed the progress
of the most significant PET improvements made in the most recent 17
years, see "400+ times improved PET efficiency for lower-dose
radiation, lower-cost cancer screening," 3D-Computing, Jun. 30,
20010, ISBN: 0970289707, which is incorporated herein by reference
in its entirety. During that time period, the efficiency of PET
improved at a rate of between two and three times every five years.
The analysis included technical literature, patents (including
those assigned to GE and Siemens) and also PETs that were built as
prototypes at several universities but were never commercialized.
At the current improvement rate of PET advancement, it would
conservatively take several decades of improvements for the
radiation dose necessary for a PET procedure to come within the
maximum radiation dose recommended by the ICRP.
[0018] What is needed is a means for increasing the accuracy and
efficiencies of PET devices enabling caregivers to more accurately
diagnose aliments related to the functionality of body organs and
not just inferences from the structure of the organs. Additionally,
what is needed is a quantum advance forward in PET devices and
procedures wherein patients can receive the benefits of PET imaging
without the associative risks from the radioactive doses necessary
for the procedures. Finally, what is needed is a means for reducing
the associated risks and increasing detection efficiencies
associated with PET imaging procedures to such an extent that the
benefits of PET imaging can be applied in well-body care and
preventative medicine strategies for apparently healthy individuals
as a standard health assessment and diagnostic tool for regular,
periodic checkups.
SUMMARY OF THE INVENTION
[0019] The present invention is directed to a system, method and
software product for increasing the efficiency of a PET device. The
present invention is directed to a series of improvements which are
concatenated and relates to provide the efficiencies of over 400
times that of prior art PET devices.
[0020] A modular, digital system, fully programmable and scalable
for a multi-modality, open (to accommodate claustrophobic or
overweight patients, with the option of closing the detector, to
increase efficiency), utilizing both Positron Emission Tomography
(PET) and Computed Tomography (CT) in one unit is presented herein
for VME and IBM-PC based platforms. The present invention fully
exploits the double photon emission occurrence and allows for
annual whole-body screening for cancer and other systemic
anomalies; only 1/30 the radiation dosage; a reduction in scan time
to 4 minutes for an axial Field of View (FOV) of 137.4 cm as
opposed to 55 minutes for an axial FOV of 16 cm; a decrease in
examination cost by 90%; an increase in sensitivity, providing
physicians with additional clinical information on a specific organ
or area and contribute to the specificity in detecting and
assessing cancer.
[0021] The present system collects digital data from multiple
electronic channels. Each electronic channel carries the
information (64-bit) of all sensors included in a given view angle
of the detector. The 64-bits data packets acquired at 20 MHz by
each channel with zero dead-time are correlated with neighboring
information and processed in real time by a DSP processor to
improve the signal-to-noise ratio and extract and measure particle
properties, resulting in the identification of the particle's
position, accurate energy measurement, Depth of Current PET devices
Interaction (DOI), and the timing measurements. A thorough
real-time algorithm that best identifies the photons can be
executed because the 3D-Flow sequentially-implemented, parallel
architecture (SIPA) allows for processing time to be extended in a
pipeline stage beyond the time interval between two consecutive
input data by configuring by-pass switches in parallel with the
processors. Very low power consumption drivers drive short,
equal-length PCB traces between 3D-Flow chips, solving the problem
of signal skew, ground bounce, cross-talk and noise. The
electronics validate and separate events from the different
modalities (PET/CT); PET events are checked for coincidences using
a circuit sensitive to radiation activity rather than the number of
detector elements. Both PET and CT examinations occur at the same
time in a stationary bed position using a detector with a long
axial FOV, avoiding motion artifacts, increasing throughput,
reducing examination cost, reducing radiation to patients,
increasing resolution, improving data quality, and reducing
erroneous readings (false positives). The saturation of the
electronics in prior art PET is overcome by using a system with an
input bandwidth of 35 billion events per second distributed over
1,792 channels. The output bandwidth is selectable to sustain the
activity generated by the maximum radiation that a PET/CT should
ever receive.
[0022] The pipe-line architecture of the present invention runs
contrary to that known in the prior art. Rather than a task being
divided into incremental sub-tasks for execution of each processor
in a pipeline circuit, the entire task is accomplished at a
pipeline processor prior to the data moving out. When a data enters
a unit, it will stay there until the entire task is completed. The
result will then "walk," one step at a time, through to the exit
(stopping for one cycle at each register at each unit but without
being further processed).
[0023] Also disclosed is a detector assembly capable of determining
extremely accurate DOI measurements. A detector element assembly is
coupled to a photomultiplier (PMT) at one end and to an Avalanche
Photodiode (APD) at the other end. The APD size is typically
smaller than the PMT (and crystal) size, thus a light guide convoys
the light from the larger surface coupled to the crystal detector
to the smaller area of the sensitive APD. The crystal is made of a
single continuous block of material, or it can be made of two
sections. One section is coupled to the PMT is a continuous
(single) block of crystal. The entire detector (barrel or a section
covering a large portion of the human body) can be made of a single
piece of crystal which is then coupled to several PMTs. A second
section is coupled to a previous crystal block in one side and is
coupled to the light guide on the other side. This can be made of
pixel (1.times.1 mm to 5.times.5 mm in size). A reflective material
is placed between pixels in order to reduce on adjacent PMTs the
spread of the light originated by the interaction of the incident
photon with the crystal. The assembly of the detector provides the
possibility to change the thickness of the entire crystal and the
percentage of the thickness of the crystal with equal length
reflecting slits compared to the solid crystal. The typical
function of the PMT is to accurately measure the photon arrival
time, its energy, and spatial resolution, while the function of the
APD can be less important and just for a simpler function of
providing the energy information on the other side of the crystal
that would allow the calculation of the depth of Interaction (DOI).
The purpose of the cut (slits) between small crystals (pixels) is
to reduce the number of photomultipliers affected by the light
generated by an event (or interaction between the incident photon
and the crystal). The length of these cuts which separate two
crystals has to be determined experimentally and is different from
crystal to crystal. The optimal solution will be when the highest
spatial resolution, low detector dead time, and good separation of
pileup events is achieved. The optimal solution is determined by
changing the centroid calculation and the pileup separation
real-time algorithm, together with the change on the length of the
slits. For some fast crystals, the cut of the crystals (slits) is
not necessary.
[0024] Unlike prior art PET electronics which are typically
implemented asynchronously, the electronics of the present
invention is synchronous. The analog-to-digital converters can
sample the signals from the PMT or APD synchronously or
asynchronously. Typically, detectors with long decay time are
sampled synchronously at a higher rate (two to five times faster
than the decay time), while fast crystals can be sampled
asynchronously with a fixed delay from when a trigger generated
from the constant fraction discriminator occurs. In any event,
regardless of the technique used, a constant fraction discriminator
triggers on the photon's arrival time. This trigger signal is sent
to a time-to-digital converter, which measures the photon's arrival
time (with respect to the system clock).
[0025] The electronics in the 3D-Flow DSP photon detection board
(the data reduction stack) can provide accurate information on all
the above parameters (time-stamp, total energy, and DOI
measurements) because each channel has a dedicated set of DSP
processors. Those DSP can compute complex calculations on each
signal that arrives from the PMT and can correlate that signal with
the eight neighboring signals. Each has highly programmable
computing capability and neighboring (eight, twenty-four, etc.)
data exchange, allowing for the extraction of highly accurate
spatial resolution information on the interaction between the
photon and the crystal. The programmability of the present
processors and its architecture allows the execution of any
algorithm (i.e., any DOI measurement with any technique), even if
it takes more time than the interval between any consecutive input
data. The present invention can measure more accurately the total
energy by summing the energy of the eight neighbors (rather than
only three neighbors as implemented in the current PET). It also
utilizes a narrower energy window for better separation of the
scatter events from the good events and thereby achieves the goal
of "rejecting more scattered events than good ones."
[0026] Higher efficiency and greater accuracy (image resolution)
are made possible by summing eight neighbors or more with the head
of a cluster, the calculation of the time-stamp, the execution with
zero dead-time of a complex programmable real-time algorithm for a
time longer than the time interval between two consecutive input
data, and offers the possibility of extending the FOV in a
cost-effective manner to capture a greater number of photons.
[0027] The calculation of the DOI on any of the three detector
implementation techniques, and the centroid calculation based on
the information from the eight neighbors on four sides preserves
and increases the spatial resolution compared to the prior art
"Anger Logic" technique which is based on information from just
three neighbors from on only two sides.
[0028] The "time coincidence detection board," in addition to
complementing the features of the "the photon detection board,"
increases the sensitivity of the PET by accepting through the LVDS
serial input/output lines a string of 64-bit information relative
to the photon found (or any bit string such as the standard PET
link format). The information of the 64-bit string is specified in
the proposal (page 18) and can be received from the proposed
"3D-Flow DSP photon detection board" or from any other board
providing information on the photon through LVDS links. The "time
coincidence detection board" routes the data from several LVDS
input lines to fewer output lines. Events with the same time-stamp
are sorted and compared to different groups of detectors to find
coincidences.
[0029] In accordance with an exemplary embodiment of the present
invention, with another aspect of the present invention, the PET
improves the energy resolution by calculating the total energy of
the incident photon, even when the photon strikes a crystal coupled
with the boundary of two PMT or APD, by using its capability to
exchange data with eight (or twenty-four) neighbors with no effect
of detector boundaries. Superior spatial resolution is achieved by
calculating the "X" and "Y" position of the incident photon based
on the information of all eight neighboring PMTs (or APDs) with
respect to any PMT (or APD) element where the local maxima was
found (instead of only three neighbors or limited by detector
segmentation with boundaries as is implemented in the current PET).
Additionally, a photon's arrival time is detected and used for
assigning a time-stamp to each event. This arrival time is then
examined for time coincidence with any acceptable opposite detector
element that received a hit. Finally, the present invention
provides the possibility of executing complex real-time algorithms
(e.g., calculating DOI measurements based on different detector
implementations) on high-rate incoming data by using the massively
parallel 3D-Flow architecture.
BRIEF DESCRIPTION OF THE DRAWINGS
[0030] The novel features believed characteristic of the present
invention are set forth in the appended claims. The invention
itself, however, as well as a preferred mode of use, further
objectives and advantages thereof, will be best understood by
reference to the following detailed description of an illustrative
embodiment when read in conjunction with the accompanying drawings
wherein:
[0031] FIG. 1 is a flowchart depicting the steps necessary for
increasing the photon capture efficiency of a prior art PET to that
disclosed in exemplary embodiments of the present invention;
[0032] FIGS. 2A-2D are diagrammatic comparisons of the relationship
between the increasing FOV and Lines of Response (LORs) in
accordance with exemplary embodiments of the present invention;
[0033] FIG. 3A is a diagram of a prior art detector crystal
optically coupled to a 2.times.2 PMT module;
[0034] FIG. 3B is a diagram of a detector crystal optically coupled
to a PMT in accordance with an exemplary embodiment of the present
invention;
[0035] FIG. 4 is a diagram of a digital signal processor for
implementation in current PET systems in accordance with an
exemplary embodiment of the present invention;
[0036] FIG. 5 depicts the flow of results (photons identified by
the real-time algorithm in the 3D-Flow stack) from the data
reduction stack to the coincidence circuit in accordance with an
exemplary embodiment of the present invention;
[0037] FIG. 7 depicts a coincidence detection scheme in which only
those candidates found within a time of 50 ns are compared (no more
than four are expected); the candidates from different detector
blocks may require different numbers of clock cycles to reach the
exit point, thus a sorting/resynchronizing circular buffer realigns
the events in the original sequential order and within a fixed
delay time from when they occurred in accordance with an exemplary
embodiment of the present invention.
[0038] FIG. 7 graphically depicts a circuit which requires only six
comparisons amongst four photons (A-B, A-C, A-D, B-C, B-D, and C-D)
every sampling period of the signals from the detector in
accordance with an exemplary embodiment of the present
invention;
[0039] FIG. 8 is a flowchart depicting the sequence of operations
for the implementation of the circular buffer for sorting and
regaining fixed latency of events in accordance with an exemplary
embodiment of the present invention;
[0040] FIGS. 9A-9C depict a scintillation detector assembly as is
well known in the prior art;
[0041] FIG. 10 is a diagram of a detector assembly having two
sensors for measuring the depth of interaction to correct the
parallax error in accordance with an exemplary embodiment of the
present invention; and
[0042] FIGS. 11A-11B depicts a scintillation detector assembly
having a sensor on either end of the detector which absorbs a
photon in accordance with an exemplary embodiment of the present
invention.
[0043] Other features of the present invention will be apparent
from the accompanying drawings and from the following detailed
description.
DETAILED DESCRIPTION OF THE INVENTION
[0044] The present invention, referred to internally herein as the
three-dimensional complete body screening (3D-CBS) compared to the
current Positron Emission Tomography (PET), encompasses a plurality
of inventions disclosed herein and in related patents and
co-pending patent applications identified throughout this
disclosure. The scope of the corpus of inventions that comprises
the 3D-CBS may not be fully appreciated without carefully examining
the 3D-CBS from various perspectives which are important for
medical professions that engage in human body scanning. Therefore,
prior to discussing the exemplary embodiments of the present
invention, the differences between the 3D-CBS system of inventions
will be discussed with respect to a PET known in the prior art.
Here, it should be understood that, although the 3D-CBS system will
be discussed with regard to implementation in an exemplary
embodiment of a PET, those of ordinary skill in the art will
appreciate that the disclosed inventions are readily applicable to
various types of tomography, such as Computerized Axial Tomography
(CAT or CT), Single Photon Emission Computerized Tomography (SPECT)
and PET CT. The following discussion examines the present 3D-CBS
system with regard to increasing PET efficiency using the detection
of coincident photon pairs as a metric, increasing image resolution
and finally increasing patient usability.
[0045] FIG. 1 is a flowchart which illustrate an exemplary method
for implementing the exemplary embodiments of the present invention
on a prior art PET with regard to specific objectives. In
particular, FIG. 1 is a flowchart depicting the steps necessary for
increasing the photon capture efficiency of a prior art PET device
to that disclosed in exemplary embodiments of the present
invention. The PET efficiency described in the flowchart on FIG. 1
is further conditioned on all captured photons being identified as
one pair of a coincident pair of photons (i.e., coincidence
detection).
[0046] It is generally accepted by those practicing in the relevant
art that primary source of poor PET efficiency resulting from lost
photons results from inefficiencies in crystal detectors. While
detector crystals do not have perfect stopping power and do not
capture every photon in range, as measured by the industry and
independent researchers, the operating efficiency of detector
crystals has been demonstrated to be 80% to 95%. Thus, according to
the industry, 80% to 95% of the photon incidences at a detector
crystal are converted into electrical signals. By contrast, the
inventor of the present invention has independently discovered that
the efficiency of prior art PET electronic can be calculated at
approximately 8% (discussed in greater detail below). Inefficient
PET electronics is partially due to dead-time resulting from
bottleneck (e.g., multiplexing of data from many lines to a single
line, saturation on input, processing, saturation on output)
present at any stage of the electronics. Another shortcoming of
prior art PET electronics is saturation of the electronics at the
output stage due to the limiting architecture of the coincidence
detection circuitry. These and other shortcomings of the prior art
have been overcome and the efficiency of PET devices improved by
using a special massively parallel-processing system architecture
with digital signal processing on each electronic channel in
accordance with an exemplary embodiment of the present invention
(step 102). The presently described processing system architecture
is capability of fully processing all data captured (no electronic
system dead time), without saturating the electronic system and
further has data exchange capability between neighboring
processors. The presently described processing system architecture
allows for the detection of more photons, more accurately.
Moreover, by implementing the presently described processing system
architecture and overcoming the inherent inefficiencies of the
prior art, the architecture e allows for the detection of more
photons and or the implementation of a simplified, more efficient
coincidence detection circuit. The present architecture is
described in greater detail below with respect to FIGS. 4-8.
[0047] Furthermore, the presently described processing system
architecture allows for the implementation of a simplified detector
assembly design for eliminating boundaries between detector
elements (step 104). Additionally, boundaries between electronic
channels are likewise eliminated because digital signal processors
associated with each electronic channel have the capability to
communicate with neighboring processors. Because the boundary is
eliminated, a detector may share light with other detector
crystals, which is converted into a signal by its transducer, and a
processor can compare its channel signal with each of its neighbors
for more reliable identification of photons. In any case, each
electronic channel exchanges its data with all its neighbors over
the entire detector. PET inefficiencies due to boundary limitations
and their solutions are discussed below with regard to FIGS. 3A and
3B.
[0048] Additional improvements in PET efficiency are realized by
executing complex real-time algorithms on the digital signal
processors of each electronic channel in accordance with an
exemplary embodiment of the present invention (step 106). The
parallel-processing architecture and the improved and simplified
detector assembly enable the execution of these algorithms, which,
among other advantages, allow for the detection of more photons
more accurately.
[0049] The parallel-processing architecture and the improved and
simplified detector assembly increases the processing bandwidth
making it possible to efficiently and accurately handle additional
signals. Therefore, the PET device can be modified for captured
additional photons, such as by increasing the Field of View (FOV)
and/or length of the detector in a cost-effective manner
(permitting to use also economical crystals) (step 108). Typically,
the FOV of a prior art PET device is in the range of sixteen
centimeters (16 cm). This equates to an efficiency for prior art
PET devices in human scanning to approximately 0.02% at best
because the radiation in the patient is in areas of the patient
that are outside the FOV of a prior art PET device. The advantages
of increasing the FOV are discussed below with regard to FIGS.
2A-2D. One result of the increasing the FOV is that the solid angle
increases correspondingly, which in turns it allows the capture
more photons within the FOV (step 110).
[0050] FIGS. 4A-4D are diagrams of a digital signal processing for
implementation in current PET systems in accordance with an
exemplary embodiment of the present invention. The design of
circuit 400 is flexible enough to be used in several models of PET
devices manufactured by various manufacturers.
[0051] In accordance with one exemplary embodiment of the present
invention, the design specification of DAQ circuit 402 is as
follows: [0052] 16 digital input channels (16-bit word-wide per
channel); [0053] Two input clocks at 20 MHz and 40 MHz with
internal; [0054] PLL on each FPGA chip that provides the internal
timing at 320 MHz; [0055] Two differential lines for output results
(LVDS); [0056] Time-to-digital converter measuring photon's arrival
time on each channel with resolution of 500 ps; [0057] Capability
to execute in a programmable form, complex real-time algorithms
with an execution time longer than the time interval between two
consecutive input data. For instance, photon-detection algorithm,
DOI measurements in PET or particle detection in HEP applications,
or any real-time processing (graphic processing, data compression,
etc.). [0058] Capability of fast data exchange with neighboring
3D-Flow.TM. processors (North East, West, and South), which allows
the correlation of signals that were split between several
channels. This allows also clustering and local maxima calculation.
[0059] Capability to trigger on any channel that has been acquired
and processed in parallel on all channels with zero dead time.
[0060] Capability to funnel results from 16 input channels to one
(or two) output channels via routing algorithms executed on 20
processors 3D-Flow-pyramid accommodated in 5 FPGA chips. [0061]
Four serial I/O interfaces for 3D-Flow.TM. program loading,
initialization, and system monitoring during data taking PCI
interface; [0062] The testability with: a) JTAG chain through the
29 large components; b) 70 LED; c) 120 test points at a 120-pin
connector; and d) 50 test points scattered at different locations
on the board which permit monitoring/debugging of critical
functions/timing; [0063] The board is designed to work: a)
`stand-alone` to process data at a high rate; b) in a system made
of several boards controlled only by RS232; or c) stand-alone or in
a system controlled by PCI interface; [0064] The board is designed
and implemented in such a way that any clock pin of any 3D-Flow.TM.
FPGA chip in any board of the system (even when the boards belong
to different crates or chassis) will not have a skew with any other
3D-Flow.TM. FPGA clock pin that will exceed 40 ps; and [0065] The
high parallelism of the internal units of the 3D-Flow.TM. processor
also allows the execution of complex real-time algorithms.
[0066] Each of processors 410 in one layer of the 3D-Flow stack 422
(see FIG. 4C) executes in parallel the real-time algorithm, from
beginning to end, on data received from the PET detector, while
processors at different layers of 3D-Flow stack 422 operate from
beginning to end on different sets of data received from the PET
detector. The present system architecture consists of several
processors arranged in two orthogonal axes: one layer is an array
of 3D-Flow processors 410, where each processor is interconnected
to its four neighbors through North, East, West and South ports
(see FIG. 4B). Several layers, assembled one adjacent to another to
make a system, is called a "stack," represented in FIG. 4A as stack
422 which is responsible for photon detection and data reduction.
The first layer is connected to the input sensors, while the last
layer produces the results processed by all layers (layer A-D) in
the stack, with the out-results sequenced in the precise order of
the in-data from the input sensors (see FIG. 4D). Data and results
flow through the stack from the sensors to the last layer. An
electronic channel consists of one set of processors 410 connected
from the bottom port of one chip to the top port of an adjacent
chip (with the top port of the first chip connected to the signal
received from the detector and the bottom port of the last chip
connected to 3-D Flow pyramid 424). 3-D Flow pyramid 424 comprises
coincidence circuitry through a pyramidal funneling structure of
processor vertices.
[0067] The 3D-Flow architecture extends the execution time in a
pipeline stage beyond the time interval between two consecutive
input data using the bypass concept described in U.S. application
'207, discussed immediately below.
[0068] Rather than requiring an ultra fast, expensive technology
capable of executing several special instructions (e.g., data
moving and data processing such as the 26 operations of the
3D-Flow) per second, or simplifying the real-time algorithm to the
point that measurements such as energy, centroid, or DOI are not
accurate, the 3D-Flow.TM. architecture permits the execution of
complex algorithms and sustains a high input data rate using any
technology (FPGA or ASIC at 0.25 micron or smaller for enhanced
performance at a higher cost).
[0069] The extension by the 3D-Flow architecture of the execution
time in a pipeline stage beyond the time interval between two
consecutive input data is illustrated by the following example: an
identical circuit (3D-Flow processor 410) is copied four times as
shown in FIG. 4C. (The number of times the circuit is copied
corresponds to the ratio between the algorithm execution time and
the time interval between two consecutive input data.) A bypass
switch coupled to each processor in each 3D-Flow in layer A sends
one data packet to its processor and passes three input data
packets and one output result from its processor along to the next
layer. The bypass switches on the 3D-Flow processors at layer B
send two input data packets along to the next layer, one output
result received from layer A and one result from its processor, and
so on. Only the processors at layer A are connected to the PET
detector and these receive only input data. The processors at layer
D send out only results. This architecture simplifies the
connection in a parallel processing system and does not require a
high fan-out from the detector electronics to send data to
different processors of a parallel-processing system. All
connections are point-to-point with several advantages in low power
consumption, signal integrity, etc.
[0070] As discussed above, in order to understand what functions
need improvement in order to increase PET efficiency, there must be
an understanding of where photons are lost in a prior art PET
device. Consequently, it should be appreciated that the lack of
efficiency in prior art PET devices is not due to inefficiency in
crystals, as has been believed in the past, but rather it was due
to the inefficiency of the electronics, which also limits the
detector assembly and the implementation of an efficient and
accurate real-time photon-detection algorithm. The solution to
overcome the inefficiency of current PET is a massively
parallel-processing system at the front-end electronics of the PET
device such as the one described immediately above, in which the
present parallel-processing architecture can be implemented in FPGA
or ASIC. Unlike other parallel-processing systems, the present
invention allows for the execution of a programmable
digital-processing algorithm on each electronic channel with
neighboring-signal correlation. Additionally, the present circuitry
can trigger on any electronic channel based on the shape of the
pulse received or based on the information from a cluster of pulses
from several neighboring elements centered on the highest pulse (or
local maxima). Thus, it can accurately measure incident photon
energy by summing 9, 16, or 25 elements, eliminating scattered
events and separating events from the different modalities
(PET/CT). It can accurately measure the spatial resolution by
interpolating the value of the sum of three (or more) elements to
the left of the local maxima and three (or more) elements to the
right for both the X and Y positions. The high parallelism of the
internal units of processor 410 allows for the execution of complex
real-time algorithms to accurately measure DOI and eliminate
parallax error of oblique photons. An oblique penetration of an
incident photon into a crystal generates a parallax error if the
depth of interaction (DOI) is not measured.
[0071] In accordance with an exemplary embodiment of the present
invention, the first bottleneck described above is overcome by
individually sampling each of the 1,344 channels at a rate of 20
MHz for a 64-bit word, sustainable continuously on all detectors
using the massively parallel-processing system at the front-end
electronics described above. A real-time algorithm that thoroughly
checks all parameters characterizing a photon is executed on the
data of an entire event and each channel is investigated to
determine if it could be the head of a cluster (corresponding to
the location of the incident photon). Furthermore, in accordance
with another exemplary embodiment of the present invention, the
processing time in one pipeline stage is extended using a series of
bypass switches, which allows for the execution of real-time
algorithms longer than the time interval between two consecutive
input data (see specifically U.S. application '207). On the
occasion where, for reasons other than the electronics (e.g., using
as crystal slow decay time) where the rate of 20 MHz cannot be
sustained for such, having the process flow handling each single
channel of the 1,344 channels means that only one channel out of
1,344 (and not one out of 56 as is in the current PET) will be dead
for the duration of the decay process in the crystal.
[0072] The second limitation of prior art electronics involves
identifying photons in time coincidence. With regard to the prior
art, a second bottleneck (in addition to the incoming data
bottleneck) occurs in the coincidence electronics because prior art
PET devices cannot handle a large number of acquisition channels,
and therefore the number of channels is arbitrarily reduced to 56
channels. The reduction is based on a simple check to find out if a
signal received from the sum of four channels is within a certain
energy window.
[0073] The limitations in detecting coincidences of the prior art
PET devices is brought about as a consequence of requiring the
electronics to compare all pairs of signals from crystals which are
points on a line passing through the patient's body. Using this
approach, for a system with n channels, all possible comparisons
(all Lines-Of-Response (LORs) of a PET) between all channels are:
(n.times.(n-1)) divided by 2 (since only the crystals which are a
point on a line passing through the patient's the result is further
divided by 2. A prior art PET device with 56 modules must then
perform about 700 comparisons along all LOR passing through the
patient's body ((56.times.55)/2(2)=720. Moreover, by increasing the
FOV from 15 cm to 140 cm, the number of 137 cm the number
comparisons along all LOR passing through the patient's body is
greatly increased.
[0074] In accordance with still another exemplary embodiment of the
present invention, the ONLY photons that are compared are those
whose characteristics show them to be a candidate for coincidence
rather than comparing all LOR used in the current PET. To
accomplish this, additional information relating to the photon is
gathered upon detection (i.e., a time stamp), and affixed to the
data signal upon exiting the data reduction stack (photon detection
stack). Using the time-stamp for finding coincidences in a PET
system is to identify all possible candidates within a
predetermined sampling time, for example, 50 ns (no more than four
candidates are expected for a radioactive dose of 5 mCi delivered
to the patient and therefore only six comparisons are made for a
coincidence only among those candidates). It is not necessary to
test all LOR as is done by the prior art, but instead move fewer
photon candidates for coincidence (less than four) to a coincidence
circuit through a pyramidal funnelling structure.
[0075] An exemplary channel-reduction and time-coincidence board is
disclosed herein for high-efficiency detection of photons in time
coincidence in PET devices. The board comprises twenty processors
(processor 410), each capable of executing up to 26 operations in a
single cycle. These processors function identically when configured
in the pyramidal funneling structure as when configured in the
photon detection stack (2-D flow stack). Each processor can execute
programmable real-time algorithms that route messages from the
parallel Top processor input port or North, East, West, South LVDS
serial input ports to the parallel Bottom processor output port or
North, East, West, South LVDS serial output ports and can execute
sorting and coincidence-detection algorithms. The board has a
memory buffer (up to 512 MBytes) to store the attenuation
correction coefficients and for de-randomizing and buffering data
flow. It has 32 pairs of LVDS differential inputs and two pairs of
LVDS differential outputs. Several transmission protocols can be
implemented, including PETLINK protocol. Two in-phase clocks at 20
MHz and 40 MHz (with PLL.times.8=320 MHz internal clock) are
distributed so as to limit the maximum skew between the clock of
any processor in the system to less than 40 ps. The circuits are
implemented in FPGA, and full programmability is dependent only on
the real-time algorithms downloaded into the processor program
memory. The board is suitable for the current PET with different
detector types and for the 3D-CBS for best PET efficiency
improvement. Typically, sixteen 3D-Flow DAQ boards are interfaced
to one coincidence board.
[0076] The original sequences of the events as they were acquired
by the detector, as well as their latency time from a location in a
layer of the pyramid (funneling section of a 3D-Flow processing
system) with respect to the time when they were created, are lost
at the last stage of the pyramid (vertex). The reason is that
events have followed different paths (short and long) when moved
through the pyramid (see FIG. 5).
[0077] The task of this stage (or vertex of the pyramid), which is
implemented with a layer of 3D-Flow processors, is that of sorting
the events in their original sequence (see sequence of operations
in FIG. 8 and regaining the fixed latency time between data at
different stages. FIG. 8 is a flowchart depicting the sequence of
operations for the implementation of the circular buffer for
sorting and regaining fixed latency of events in accordance with an
exemplary embodiment of the present invention. The process begins
with the determination that data are available at a port (step
802). The data string is read from the port prior to fetching data
from another port (step 804). Finally, the "Time ID" is extracted
from the data and the "write-pointer" of the "circular buffer"
calculated from the "Time ID." The "write-pointer window" is
incremented, data is read from the "circular buffer" and
"write-back zero," the "read-pointer" is incremented and the read
data is sent to the three selected out-ports during the same
time-slot of 60 ns. sequence of operations implementing of the
circular buffer for sorting and regaining fixed latency of
events.
[0078] FIG. 5 depicts the flow of results (photons identified by
the real-time algorithm in the 3D-Flow stack) from the 3D-Flow
stack to the coincidence circuit. (A stack is the section of the
circuit where single photons with 511 KeV are detected.) The right
side of the figure shows the flow of results from one stage of the
3D-Flow system to the next stage with the relation of the time
delay of the data in different stages. The real-time algorithm and
its implementation, with the 3D-Flow providing the results, is
shown on the top left of FIG. 5 as output from the 3D-Flow DAQ
stack and is also described in more detail in co-pending U.S.
applications '904 and '532.
[0079] The circular buffer memory in the center of the figure
receives the data from the last layer of the pyramid. The program
loaded into the 3D-Flow processor implementing the circular buffer
reads the field of the time-stamp of the event received from the
pyramid and uses the value of its content to calculate the address
(write-pointer) of the circular buffer where the event just
received should be stored. This operation has the effect of sorting
and regaining the fixed latency delay between data. At the system
speed, the circular buffer is read out when all photons with a
given time stamp have been stored in the circular buffer. The
reading of the circular buffer(s) at any given time will provide
all photons that occurred `n` time periods before in the
detector.
[0080] There are several ways of using the scheme of the circular
buffer described above for detecting all possible photons belonging
to a specific time period `n.` One simple example is described
herein, while an example for a more general application requiring
maximum photon detection with the possibility of increasing the
output bandwidth of the system is described further below. In order
to find a coincidence, a signal from a detector block needs to be
compared with the signal from another detector block. For the sake
of convenience, the detector blocks are grouped in sectors, and
only four sectors are defined in this example. All detector
elements connected by lines that do not pass through the patient's
body are grouped together in a sector (see top right part of FIG.
6). This scheme requires the implementation of four circuits of the
type shown in FIG. 5. In accordance with an exemplary embodiment of
the present invention, with one exemplary embodiment for an
implementation the system comprises 1,152 separate channels. For
each sampling time period, the single photon detected in each of
the sectors will be compared with the photon detected in the other
sectors in the 3D-Flow processors of 158 in FIG. 6. (In the very
unlikely case that more than one photon is detected, the memory
cell of that location is overwritten and the last value written is
the one that will be compared).
[0081] Only those candidates found within a time of 50 ns are
compared (no more than four are expected, resulting in six
comparisons being made). The candidates from different detector
blocks may require different numbers of clock cycles to reach the
exit point; thus, a sorting/resynchronizing circular buffer
realigns the events in the original sequential order and within a
fixed delay time from when they occurred. The left portion of FIG.
6 shows how many types of 3D-Flow components are required to
implement the different functions. FIG. 7 shows the circuit, which
requires only six comparisons amongst four photons (A-B, A-C, A-D,
B-C, B-D, and C-D) every sampling period of the signals from the
detector. This technique is advantageous compared to approximately
700 comparisons every 250 ns, for the prior art PET discussed
above, and provides a rate of coincidences found up to 40 million
coincidences per second instead of 4 million coincidences per
second, as is the limitation of prior art PET devices.
[0082] The following is a general scheme, based on the requirements
of the maximum radiation dose delivered to the patient and the
complexity of the coincidence-detection algorithm for implementing
the circuits at the output of the 3D-Flow pyramid for sorting the
photons in the original sequence, regaining a fixed latency time
with respect to when the event occurred in the detector, and for
identifying all coincidences. The basic idea of the approach is
very simple. There is no segmentation of the detector in sectors as
has been done heretofore. If the radiation delivered to the patient
creates 80.times.106 single photons per second, the circuit
described above for sorting and realigning the latency needs to run
also at 80.times.106. A single circular buffer is implemented at
the speed equal to or higher than the rate of the single photon
created. Each photon detected within the sampling time window of 50
ns is compared with all other photons of the same time window
(e.g., six comparisons for four photons, ten for five photons,
fifteen for six photons, or (n.times.(n-1))/2), regardless of
whether or not the x, y position of the two photons being compared
lie along a line passing through the patient's body. A 3D-Flow
processor can be used for implementing the comparison circuit. A
set of 3D-Flow processors is working in parallel to perform all
comparisons of detecting coincidences within a sampling period. For
example, one 3D-Flow chip is sufficient for a 5 mCi dose to the
patient corresponding to about 80.times.106 single photons per
second activity of a PET with about 150 cm FOV. The number of
comparisons are much fewer, compared to the approach used in the
prior art PET devices.
[0083] The format of the sequence of bits sent out from the 3D-Flow
time coincidence detection and buffer board can be the standard
PET-Link format, or it can provide additional information, such as
the time of flight and depth of interaction in order to allow the
image reconstruction software on the workstation to build better
images.
[0084] The format of the output word of the "coincidences" (pair of
photons) from the 3D-Flow pyramid, time coincidence detection and
buffer memory board suggested in co-pending U.S. applications '904
and '532 is the following:
[0085] bits 0-19 crystal spatial ID (hit1);
[0086] bits 20-23 Depth of inter. (hit1);
[0087] bits 24-29 photon energy (hit1);
[0088] bits 30-33 time-of-flight (hit1 and hit2);
[0089] bits 34-53 crystal spatial ID (hit2);
[0090] bits 54-57 Depth of inter. (hit2); and
[0091] bits 58-63 photon energy (hit2).
[0092] Two 20-bit fields for spatial ID of hit1 and hit2 allows for
coding up to 1,048,575 crystals. Two-4-bit DOI fields allow for a
depth of interaction of both hits with up to 1.56 mm resolution
when crystals of 25 mm thickness are used. The energy of the two
photons is coded in 64 intervals from the smallest to the highest
energy value. The 4-bit field for the time of flight makes it
possible to locate, within 7.5 cm resolution, the point of
interaction along the line which connects the two crystals, and to
measure up to 75 cm the distance in any direction inside the
patient's body. The maximum measurement could be increased by
changing the coincidence time window parameter. For instance, a
3-ns coincidence time window parameter will allow the measurement
of any interaction that had traveled up to about 90 cm inside the
patient's body. The coincidence board receives the data relative to
the photons validated by the real-time algorithm executed on the
3D-Flow DAQ boards. It then performs the functionality attenuation
correction described in U.S. applications '904 and '532, separating
the photons found into the two modalities (PET and CT), the channel
reduction and the coincidence identification. The board stores the
results and the coincidences found (or the single photon validated
by the algorithm for CT when the buffer memories on the DAQ boards
are not installed). Results are read from two LVDS links or
directly from the buffer memories by a PC CPU via the PCI bus and
are sent to the graphic workstation via a standard high-speed local
area network.
[0093] The technique of detecting pairs of photons in time
coincidence, as described in this document and in co-pending U.S.
applications '904 and '532, offers great advantages and simplifies
the implementation of the hardware circuit. This technique is
related to the maximum radiation dose allowed to the patient. For
example, when assuming a sampling rate of the detector every 50 ns
and 80 million single photons per second being the rate of "good"
four photons, regardless of the number of detector electronic
channels (which is assumed to be about 1800 for the proposed 140 cm
FOV 3D-CBS), the number of comparisons needed will be
n*(n-1)/2=(4*3)/2=6 comparisons every 50 ns, which is equivalent to
120 million comparisons per second. This task can be easily
performed by current economical microprocessors (even by a FPGA
electronics). By contrast, if it were necessary to achieve the same
performance using the technique implemented in a prior art PET
device, the number of comparisons for the same detector with about
1800 electronic channels will be n*(n-1)/4=(1800*1799)/4=809,550
comparisons every 50 ns, which is equivalent to about 16 trillion
comparisons per second. This requirement has no practical solution.
As a compromise, current PET manufacturers do not use full
granularity of all electronic channels of the PET detector. As
discussed above, one solution offered by the prior art implements a
sampling rate of 250 ns based on a granularity of only 56
electronic channels, although the PET detector has 1344 electronic
channels.
[0094] Cost of the 3D-CBS device is reduced, or kept at least to a
minimum, through the use of low cost detector crystals. One type of
scintillator crystal known for its cost effectiveness is the
bismuth germanate (BGO) crystal. An even lower cost crystal is the
sodium iodate (NaI) crystal; however, the disadvantages associated
with NaI crystals have discouraged a large segment of the PET
industry to other more expensive crystal detectors, as mentioned
elsewhere above. NaI crystals are less dense and have less
"stopping power" of the 511 keV photons than BGO crystals. BGO is
more rugged, and allows for higher detection efficiency.
Additionally, BGO is not count-rate limited, thus practitioners are
encouraged to inject ever larger dosages of isotopes in their
patients because the BGO can, it has been surmised, detect more
counts and more counts result in clearer scans and sharper images.
In fact, some estimates place BGO crystal usage at almost ten times
that of NaI. Although the NaI crystal may have lower stopping power
than the BGO, it provides a stronger signal.
[0095] Therefore, in accordance with another exemplary embodiment
of the present invention, an improvement in the PET spatial
resolution may be achieved by means of a more accurate measurement
of the depth of interaction (DOI) using either low cost crystals
such as BGO, or the NaI crystal which has an even lower cost. The
photon's stopping power of the NaI crystal is increased by
fabricating a thicker NaI detector with a stronger signal in
proportion to a comparable BGO detector. With a renewed interest in
NaI detectors, there is a likelihood that NaI crystals will be
grown ever larger; in fact, it is technologically possible to build
a single barrel to cover the entire surface of the patient's body.
However, cost-efficiency criteria will most probably dictate an
optimal segmentation and separation of the crystal that will cover
most, but not all, of the patient's body.
[0096] Measuring the DOI is important for correcting the parallax
error. Parallax is the error that results from assuming that
photons strike the detector at 90 degrees to its face. With regard
to FIGS. 9A-9C, a scintillation detector assembly is depicted as is
well known in the prior art. The assembly comprises crystal 902,
light amplifiers 904A and 904B and corresponding detectors 906A and
96B. Crystal 902 might be any type of crystal which interacts with
a photon so as to produce a scintillation or rapid flash of light
in the interior lattice structure of the crystal. Typically,
crystal 902 is optically coupled to one or more optical amplifiers
which have a detector integrated therein. Thus, as a practical
matter, amplifiers 904A and 904B and corresponding detectors 906A
and 906B may be Photomultipliers (PMTs), Avalanche Photodiodes
(APDs) or some other type of light emitting diode; however, each
amplifier-detector combination will have a signal output (a
channel) for outputting the amplified signal to the processing
electronics.
[0097] With regard to the parallax effect, notice from FIG. 9A that
incident photon .gamma.900 is approaching crystal 902 at an oblique
penetration (instead of being perpendicular) to the face of the
crystal looking toward the emitting source. When a photon enters
the crystal at 90 degrees, its X-Y position can be easily
calculated from the detectors which perceive the scintillation
effect in the crystal, the XY position through a centroid
calculation. An exemplary centroid calculation for a 2.times.2
detector array(detectors A, B, C and D) is: X m = ( A + B ) - ( C +
D ) A + B + C + D ##EQU1## Y m = ( B + D ) - ( A + C ) A + B + C +
D ##EQU1.2##
[0098] (A better calculation for determining .DELTA..sub.x is the
ratio of the sum of the energies of all sensors at the west of the
central element, divided by the sum of all sensors at the east of
the central element (.DELTA..sub.x=.SIGMA.E.sub.W/.SIGMA.E.sub.E).
Similarly, for the calculation of .DELTA..sub.y, the ratio of the
sum of the energies of all sensors at the north of the central
element, divided by the sum of all sensors at the south of the
central element
(.DELTA..sub.y=.SIGMA.E.sub.N/.SIGMA.E.sub.S.)).
[0099] The depth at which the photon interacts with the crystal is
unimportant in this case where the photon penetrates the crystal
perpendicular to the face, because it will interact somewhere along
a line oriented in the Z direction formed by the intersection of an
X plane and a Y plane (i.e., the LOR is found perpendicular to the
X-Y planes). This presumes that all lines of response between
coincidental pairs of detectors intersect the center point of the
barrel which is very imprecise. In practice, once the detector
elements 906 A and 906 B receive an optical signal, an analog
signal is produced at output 908 and sent to the PET electronics
(the coincidence board(s)). Generally, the PET electronics compare
all of the possible LOR for coincidences, even those connecting two
detectors, that did not receive a hit. When a coincidence is
determined, the resulting LOR is used for generating the image.
However, the parallax effect shifts the placement of the endpoints
of the LOR along the Z axis to some default depth, such as the mid
point or face of the crystal. The error is apparent on FIG. 9C,
where both LOR 920 and LOR 922 are correctly spatially positioned
on the X-Y plane of detector 902, but only LOR 920 is at the proper
depth. Often, if a DOI calculation is not performed, the LOR is
found by correspondence using a default depth (e.g., midway down
the detector, on its face, etc.) The result of not calculating a
DOI are graphically illustrated in FIG. 9C by the separation
between LOR 920 and LOR 922.
[0100] Therefore, the parallax error resulting from of incident
photons with angles different from a 90-degree measurement is
corrected by determining an accurate interaction depth and using
the depth to properly place the LOR. DOI is determined by comparing
the photon's energy, as captured by two different detectors, and
relating the difference to the interaction depth of the photon in
the crystal. Best results are obtained when the detectors are
positioned to maximize variations in energy based on the depth of
interaction. One detector should offset depth, with respect to the
Z axis. In accordance with an exemplary embodiment of the present
invention, the measurement of the depth of interaction to correct
the parallax error of incident photons with angles different from
90 degrees can be performed by using two sensors, for instance
Photomultipliers (PMT) or Avalanche Photodiodes (APD) on both sides
of the detector, one internal to the barrel and the other external
to the barrel. For instance, by using an array of photomultipliers
internally and externally and then interpolating the signals
received by the two sensors. FIG. 10 is a diagram of a detector
assembly having two sensors for measuring the depth of interaction
to correct the parallax error in accordance with an exemplary
embodiment of the present invention. There, crystal 1002 is
optically coupled to amplifier-detector 1004 and light guide 1016,
which is coupled to amplifier-detector 1018. One exemplary
embodiment employs Photodiodes or APD internally, rather than a
PMT, to improve efficiency. Furthermore, the semiconductor will not
absorb or scatter many photons that penetrate the face of the
crystal because it is comprised of an extremely thin material, only
a few hundred microns. In addition, because the detector obscures
only a portion of the face of the crystal, not every photon
penetrating the crystal's face will pass through detector 1018.
Photodiodes and APD will generally cost more than PMTs and have a
lower gain, those deficiencies will probably abate somewhat as the
convenience of using Photodiodes or APD internally and externally
becomes more apparent. In the present embodiment shown in FIG. 10,
however, amplifier-sensor 1004 is depicted as a PMT, while
amplifier-sensor 1014 is illustrated as an APD. The light captured
by the two sensors, which is proportional to the energy of the
incident photon and to the distance where the photon was absorbed
by the detector with respect to the location of the two sensors, is
converted into electrical signal 1008 and 1018. The two signals are
converted into digital form, sent to the 3D-Flow processor which
computes the interpolation of the distance from the two sensors,
which is proportional to the location where the photon hit the
detector. This measurement allows for more accurately determining
the location where the photon hit the detector, thus eliminating
the parallax error, thus improving spatial resolution. Although
FIG. 10, and others, depict the detector as having been segmented
into small rectangular shapes, that depiction is not intended to
limit the scope of the present invention. Despite the fact that the
crystal detectors may be cut in small pieces, as stated above, the
entire barrel can be fabricated from several sectors,(two, four or
eight arc segments). Still further, the barrel may be constructed
as a single piece surrounding the entire body of the patient.
[0101] FIG. 10 shows the example of a detector assembly with a thin
sensor (e.g. APD) in front of the detector (side where the
radioactive source is located and the photo is hitting the
detector) and a second sensor (APD or photomultiplier) on the
opposite side of the detector. The light captured by the two
sensors interior sensor 1018 and exterior sensor 1008, which is
proportional to the energy of the incident photon and to the
distance where the photon was absorbed by the detector with respect
to the location of the two sensors, is converted into electrical
signals 1018 and 1008, respectively. Signals 1018 and 1008 are
converted into digital form, sent to the 3D-Flow processor, which
computes the interpolation of the distance from the two sensors,
which is proportional to the location where the photon hit the
detector. This measurement determines more accurately the location
where the photon hit the detector, thus eliminating the parallax
error, and improving spatial resolution. Hence, PD (APD) sensor
signal 1018 and PMT signal 1008 are linearly dependent on the depth
of interaction (Z) from the photodetector.
[0102] Here, it should be noted that, in contrast with prior art
detectors configured for DOI calculations, the 3D-CBS uses the
outputs from the exterior PMTs for the vast majority of the data to
be used for image generation. As mentioned above, the present
system is hundreds, if not thousands, of time more efficient than
the prior art PET device using only the photomultipliers.
Therefore, while the 3D-CBS architecture could easily accommodate a
complex interior sensor arrangement, such as an array of interior
sensors, there is simply no need to expend the resources on
developing interior sensors and signal channels that will be used
for only one purpose--to be compared to the exterior signals for an
interaction depth. To that end, the present interior sensors are
chosen and configured with cost effectiveness as a primary intent.
The results of the choices on the detector configuration are
strikingly different than any interior sensor arrangement hereto.
For instance, one means to achieve cost effectiveness is to reduce
the coverage area of the APD. Notice from FIG. 10 that, although
the detector 1002 has approximately the same area as the face of
PMT 1004, the coverage area of the APD is much smaller than the
face of crystal 1004. For the purposes of the present invention,
this makes absolute perfect logic. The faces of detector 1002 PMT
1004 should be comparable for better optical coupling and lowering
the risk of missing an event. The requirements for coupling APD
1014 are much less stringent. In fact, since what is sought from
APD is a reasonably accurate signal, the diode utilizes optical
guide 1014 to collect and channel the scintillation from detector
1002. In stark contrast with prior art DOI schemes, it is simply
not necessary to use the interior sensor for anything other than
collecting an optical signal to be compared with the exterior
channel signals.
[0103] Turning now to FIGS. 11A-11B, a scintillation detector
assembly having a sensor on either end of the detector is depicted
absorbing a photon in accordance with an exemplary embodiment of
the present invention. The assembly comprises crystal 1102, light
amplifiers 1104A and 704B and corresponding detectors 1106A and
1106B. Here again, crystal 1102 may be any known or heretofore
unknown type of detector which interacts with a photon to produce a
scintillation or a rapid flash of light in the interior lattice
structure of the crystal. Crystal 1102 is coupled to one or more
optical amplifier/sensors which have a detector integrated therein.
Also, as discussed with regard to FIG. 10, amplifier-sensor 1104 is
depicted as an PMT, while amplifier-sensor 1114 is illustrated as
an APD. Notice from FIG. 11B, however, that amplifier-sensor 1114
was the first to receive an optical signal from crystal 1102,
resulting in output electrical signal 1118, while at a later time
amplifier-sensor 1154 received the optical signal from crystal
1102, resulting in output electrical signal 1108. It should be
cautioned, however, that the order in which the optical signals are
received and the timing are relatively unimportant. The present
invention utilizes the energy levels at the respective sensors, not
the signal arrival times, to determine the DOI of the photon in
crystal detector 1102. The depth of interaction, not the arrival
times, is proportional to the respective signal strengths . In any
case, once electrical signals 1108 and 1118 have been generated,
they are passed to the 3D-CBS DOI electronics for integration and
depth determination. To that end, optical guide 1116 collects and
redirects the optical signal toward the active portion of APD 1114
in an extremely cost effective manner.
[0104] At present, the exterior sensors are PMTs for the reasons
discussed above. However, correction of parallax errors from
incident photons with angles different from 90 degrees can be
performed by using two sensors (Photomultipliers or Avalanche
Photodiodes APD) on both sides of the detector, one internal to the
barrel and the other external to the barrel, for instance, by using
an array of photomultipliers internally and externally and then
interpolating the signals received by the two sensors. In
accordance with one aspect of the present invention, Photodiodes or
APD is used internally that will not absorb or scatter many photons
will significantly improve efficiency of the system because of its
small thickness of material of a few hundred of microns, and a PMT
is used externally. Photodiodes or APD will cost more than PMTs and
have a lower gain; however, future technology advances will show
that it will be convenient to use Photodiodes or APD internally and
externally. Although the present invention is using an exemplary
embodiment having a detector cut (or slit) in a small rectangular
shape, the present invention is not so limited to crystal detectors
cut in small pieces. Instead, the present invention may be
implemented having a detector with the entire barrel can be made of
several sectors, four sectors, two sectors or at the limit a barrel
in a single piece surrounding the entire body of the patient. This
detector can have sensors (PMT, APD, or photodiodes) internally or
externally to the barrel.
[0105] FIGS. 2A-2D are diagrammatic comparisons of the relationship
between the increasing FOV and Lines of Response (LORs) in
accordance with exemplary embodiments of the present invention. A
PET with an axial FOV that is twice as long as the short FOV of the
prior art PET can detect four times the number of photons in time
coincidence from an organ emitting photon from the center of the
FOV. FIG. 2A and FIG. 2B assume the detector has only three rings
of detector elements. Only the LOR connecting opposite sets of
detectors within the three rings are considered instead of all
possible LORs passing through the patient's body. The top detector
elements, A, B and C, and the bottom detector elements are depicted
in the figure as elements D, E, F. For a linear source at the
center of the FOV emitting pairs of photons in time coincidence in
opposite directions, one could capture only three possible
combinations AD, BE and CF (See FIG. 2A) when SEPTA are used (septa
are lead rings between the ring-detectors that prevent photons
arriving with an angle from hitting the detector). Thus, FIG. 2A
depicts a prior art PET device with short FOV and further LOR
limiting septa.
[0106] For the purpose of understanding how the capturing of
photons is greater than double when the FOV is doubled, assume that
the representation of the detector is simplified as shown in FIG.
2B which depicts a prior art PET with the same short FOV as in FIG.
2A, but the number of photons captured increases from three to nine
when the SEPTA are removed. In the absence of SEPTA lead rings,
there are nine possible combinations of pairs of photons (AD, AE,
AF, BD, BE, BF, CD, CE, CF) which can be captured.
[0107] FIG. 2C depicts the effect of doubling the axial FOV has on
LOR. Doubling the FOV, thereby doubling the number of detector
element rings, increases the Lines of Response four times over a
prior art PET device with half the number of rings (or 12 times if
compared to 2-D mode, shown in FIG. 2A). If the FOV is doubled with
new top detector elements G, H, L, and the new bottom detector
elements M, N, P, then 36 combinations of pairs of photons emitted
in opposite directions from a linear source in the center of the
FOV are captured. The possible pairs for which a LOR could be drawn
are: AD, AE, AF, BD, BE, BF, CD, CE, CF, plus the new GM, GN, GP,
HM, HN, HP, LM, LN, LP, plus the combination of old top and new
bottom AM, AN, AP, BM, BN, BP, CM, CN, CP, plus the combination of
the new top and the old bottom GD, GE, GF, HD, HE, HF, LD, LE,
LF.
[0108] Finally, the LOR algorithm described above is extendable;
for instance, if the FOV is increased three times from that
depicted in FIG. 2B, the number of pairs of photons that can be
captured increases nine times (or 27 times if compared to the
current use of the PET in 2-D shown in FIG. 2A). If the FOV is
increased four times from that depicted in FIG. 2B, the number of
pairs of photons that can be captured increases sixteen times (or
48 times if compared to the current use of the PET in 2-D shown in
FIG. 2A).
[0109] Considering that most of the PETs (even the most advanced)
currently available in hospitals use a 2-D mode for the torso,
where only the combinations AD, BF, and CF are detected, the
difference between the prior art PET and the 3D-CBS when the FOV is
doubled, is from 3 to 36 (or 12 times). If the FOV of the prior art
PET is tripled from 16 cm to 48 cm, then the difference in captured
pairs of photons will increase 27 times when using the 3D-CBS
approach.
[0110] With reference again to FIG. 1, increasing the solid angle
also increases the photon capture efficiency by reducing the amount
of photons lost at either end of the detector barrel. Some photons
from within the detector area are also lost. Some quantity of
photons that emanate from the part of the body that is covered by
the detector leave the body at an angle that allows them to escape
the detector through the openings between the detector segments.
This quantity can be calculated as a percentage of the perimeter of
a circle drawn around the lengthwise cross section of the entire
detector not covered by the 16 cm FOV barrel.
[0111] Increasing the FOV inherently results in increasing the
solid angle and thus capturing more photons, but in addition to
FOV, decreasing the diameters of the barrel opening also limits the
solid angle. Typically, the barrel of a prior art PET device is
implemented with a constant diameter throughout and since prior art
PET devices are typically configured having the barrel's diameter
sufficiently large enough to accommodate the most robust patient
body shapes, the solid angle is high and photons are lost at the
barrel's ends. The solid angle of the present invention, on the
other hand, is limited by the FOV, but also by the diameters of the
rings of the barrel being separately adjustable for the
corresponding portion of the patient. Thus, in accordance with an
exemplary embodiment of the present invention, the rings at the
patient's head and legs may be separately configured with a much
smaller diameter than those rings corresponding to the patient's
torso, thereby greatly limiting the solid angle and reducing the
amount of photons lost at the barrel's open ends (see U.S. patent
application Ser. No. 10/453,255 (hereinafter U.S. application '255)
entitled "Gantry for Geometrically Configurable and
Non-configurable Positron Emission Tomography Detector Arrays."
[0112] With further regard to the barrel, the entire structure may
be treated as one or two cameras which process photon events
received within one group of 32.times.64 Photomultipliers (PMTs) or
two groups 32.times.32 PMTs, rather than hundreds of groups of
2.times.2 PMTs. It should be understood that, rather than a PMT, an
Avalanche Photodiodes (APDs) or some other type of light emitting
diode may be substituted; however, each amplifier-detector
combination will have a signal output (a channel) for outputting
the amplified signal to the processing electronics. This is
accomplished, in accordance with other exemplary embodiments of the
present invention, by eliminating the boundaries between crystals
and between small groups of PMTs (see U.S. applications '207, '904
and '532 and especially '024 and '255). Typically, prior art PET
devices utilize a block detector design concept in which a single
crystal is optically coupled to a 2.times.2 block (or module) of
PMTs.
[0113] FIG. 3A is a diagram of a prior art detector crystal
optically coupled to a 2.times.2 PMT module. A boundary is
established between each 2.times.2 PMT module 304 and similarly
between each crystal 302. Each 2.times.2 PMT module 304 is treated
by the PET as a small camera and photon impacts are independently
processed. For example, when photon 300 impacts crystal 302 and is
received at 2.times.2 PMT module 304, the event is processed
independently of every other 2.times.2 PMT module. If 2.times.2 PMT
module 304 cannot identify a signal as being a photon impact, the
boundary does not allow the recipient module to compare its signal
with its neighbors and that photon is lost. The identification of
the crystal of interaction in the 2.times.2 PMT block is made
through the Anger Logic shown below using only the four PMTs in the
module.
[0114] (1)
[0115] (2)
[0116] Because communication between adjacent 2.times.2 PMT modules
is impossible, centroid calculations are necessarily dependent on
the separate 2.times.2 PMT modules and lack information from
adjacent modules that is necessary for accurately determining the
point of impact for the photon.
[0117] Crystal 302 (coupled to prior art PMT module 304) is
typically subdivided into an 8.times.8 block of variable length
slits. The 8.times.8 block does not share light well with adjacent
8.times.8 crystal blocks. Moreover, edge and corner subdivisions of
each prior art 8.times.8 crystal block contribute only a small
signal compared to the contribution of the inner subdivisions of
the crystal making the identification of photon events more
difficult, and lowering the overall efficiency for the PET.
Furthermore, if a photon strikes the boundary edge between adjacent
2.times.2 PMT modules (between the edge and/or corner subdivisions
of two 8.times.8 crystal blocks), neither PMT may receive
sufficient energy to recognize the strike as a photon and the
photon is lost, further reducing the efficiency of capturing
photons for the prior art PET device.
[0118] In accordance with exemplary embodiments of the present
invention, these problems are overcome by permitting each PMT to
share and receive information (signals) with its neighbor PMT, and
further by permitting the crystal to have the same degree of light
sharing throughout (or with adjacent crystals) by using slits of
equal length (or no slits), thereby allowing sharing the light with
adjacent PMTs in all four directions with no boundaries. FIG. 3B is
a diagram of detector crystal optically coupled to a PMT in
accordance with an exemplary embodiment of the present invention.
By treating the PET as one large camera, rather than hundreds of
smaller cameras, photon impacts are more readily identified than
the prior art because there are no boundary limitation on where a
PMT may get information. Signals for photon impacts occurring on
the edges and corner blocks associated with PMT 324 are shared with
its neighbors without regard to any boundary; consequently, photon
impacts on the edges, corner and between PMTs are much more readily
identified as a photon incidence.
[0119] The point of impact of a photon's may be accurately
calculated using essentially a two-step process without regard to
boundaries in accordance with another exemplary embodiment of the
present invention. The process comprises finding a local maxima for
an impact and then calculating the precise point of impact of the
photon in a PMT cluster of a predetermined size (2.times.2,
3.times.3, 4.times.4, 5.times.5 and so on). The local maxima is
defined as the head of a cluster of PMTs (of a predetermined size
2.times.2, 3.times.3,etc.) which corresponds to the location of the
incident photon. The local maxima is found by checking the signal
(and arrival time) at a PMT with similar information in the
neighboring channels. When the local maxima is determined, the
photon's energy can be calculated by summing the energy of the
local maxima with its neighbor's energies (e.g., for a 3.times.3
PMT cluster Energy=NW+N+NE+W+C+E+SW+S+SE). The photon's precise
point of impact may then be determined by sharing light between ANY
predefined cluster of PMTs (e.g., a 3.times.3 PMT cluster), by:
[0120] (3)
[0121] (4)
[0122] Finally, and as alluded to above, perhaps the single most
cost effective area of focus for increasing PET efficiency is in
the processing electronics. Within the electrons, two primary areas
exist which limit the efficiency of the prior art PET. The first
involves limitations of prior art electronics to identify valid
photons, and the second involves limitations of prior art
electronics for identifying photons in time coincidence.
[0123] It is recognized that prior art PETs capture only 0.2
million pairs per second of the original, 1,424 million pairs of
photons per second emitted by tracer within the patient's body. It
has been further proposed that approximately 2.6 million pairs of
photons per second are remaining after the natural phenomenon of
photons being scattered or absorbed in the patient, the smaller FOV
and smaller solid of prior art PET devices, and the inherent
inefficiency of a crystal photon detector. Thus, of the 2.6 million
pairs of photons per second remaining, 2.4 million pairs of photons
are lost per second due to deficiencies in the electronics and the
detector design which accounts for the prior art PETs capturing
only 0.2 million pairs per second. Therefore, vast increases in
photon by summing 9, 16, or 25 elements. Spatial resolution can be
accurately measured on ANY cluster of 3.times.3 (or 5.times.5,
either of which are predefined) PMTs. Additionally, each processor
can execute complex real-time algorithms to accurately measure DOI
and eliminate the parallax error of oblique photons. Timing is
controlled by two in-phase clocks at 20 MHz and 40 MHz (with
PLL.times.8=320 MHz internal clock) with a skew <40 ps between
any processor clock in the system. A Time-to-Digital converter
(TDC) measures arrival time and assigns a time stamp to the photon
at each channel with 500 ps resolution. The board has 2,211
components with >20,000 pins connected with about 9,000 nets in
a PCB with only eight layers of signals and six layers for power
and ground. The board is suitable for use with prior art PETs
having different detector types and for the 3D-CBS for best PET
efficiency improvement.
[0124] The corresponding structures, materials, acts, and
equivalents of all means or step plus function elements in the
claims below are intended to include any structure, material, or
act for performing the function in combination with other claimed
elements as specifically claimed. The description of the present
invention has been presented for purposes of illustration and
description, but is not intended to be exhaustive or limited to the
invention in the form disclosed. Many modifications and variations
will be apparent to those of ordinary skill in the art without
departing from the scope and spirit of the invention. The
embodiment was chosen and described in order to best explain the
principles of the invention and the practical application, and to
enable others of ordinary skill in the art to understand the
invention for various embodiments with various modifications as are
suited to the particular use contemplated.
* * * * *
References