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W. W. Moses & J.S. Huber
Dedicated Nuclear Medical Instrumentation for Prostate Cancer
Dedicated Nuclear Medical Instrumentation for Prostate Cancer
W. W. Moses and J. S. Huber
Lawrence Berkeley National Laboratory, Berkeley, CA 94720 USA
Abstract
We describe design concepts for three nuclear medical systems optimized for imaging the
human prostate. Functional (rather than anatomic) imaging of prostate cancer can assist
in the diagnosis, treatment planning, and follow-up of therapy, and cameras optimized for
imaging a single organ / disease potentially have better performance tradeoffs than
conventional multi-purpose nuclear medicine cameras. We describe the design
considerations for a transrectal single gamma imager, a PET camera with a transrectal
probe and an external bank of detectors, and a PET camera with a pair of external curved
detector banks. Of these three concepts, the final one appears to be the most
advantageous.
Key Words: Prostate Cancer, PET, SPECT, Instrumentation
1 Introduction
Until recently, the imaging of prostate cancer has been confined to modalities that
provide anatomical images, such as ultrasound, x-ray CT, or MRI imaging. This imaging
is done to assist in several facets of the diagnosis and treatment of prostate cancer, such
as guiding biopsy, staging, planning therapy, and monitoring the effectiveness of therapy.
In the last half decade there has been progress in using nuclear medical imaging
techniques (as well as spectroscopic MRI) to image metabolic aspects associated with
prostate cancer. The emergence of radiopharmaceuticals optimized for imaging prostate
cancer has spurred interest in cameras optimized for imaging prostate cancer. This paper
explores conceptual designs for several nuclear medical imaging systems.
The American Cancer Society estimates that 232,090 new cases of prostate cancer
were diagnosed and 30,350 men died of prostate cancer in the United States in 2005.
About 90% of all prostate cancers are found in the local and regional stages [1]. Prostate
cancer suspicion is typically based on an elevated prostate-specific antigen (PSA) level or
on a suspicious node found during a digital rectal exam (DRE). More than half of all
cancers detected today are not palpable, and PSA and DRE screenings have high falsepositive rates in general clinical practice [2]. The treatment decision is mainly based on
biopsy confirmation of prostate cancer, but the diagnostic accuracy of biopsy is
problematic [3]. A new imaging technology for sensitive detection of prostate cancer is
needed to confirm initial diagnosis, guide biopsy and help guide treatment decisions. A
method to monitor response to therapy after an intervention is also needed. Current
means of assessing treatment response in prostate cancer are imprecise [4-6], so the
administration of multiple courses of a therapy is often necessary before a clear
indication of response or progression can be determined. Thus, we need a better prostate
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imaging technique to monitor response to therapy, assess the efficacy of new treatments,
and detect local recurrence sooner.
Prostate cancer imaging with radiopharmaceuticals has used three types of
instruments: (1) planar gamma cameras, (2) single photon emission tomography (SPECT)
cameras, and (3) positron emission tomography (PET) cameras. As radiopharmaceuticals
have been developed that are appropriate for use in any of these camera types, it would
seem reasonable to consider prostate-specific designs for all three of these modalities.
However, planar gamma cameras image 99mTc compounds that are designed to
accumulate in metastatic disease in the bone, and so do not accumulate in the prostate.
While they are widely used in the management of prostate cancer, these agents would
not be used in a single photon camera that is optimized for imaging the prostate. Thus,
we consider only SPECT and PET designs optimized for prostate imaging.
2 SPECT Methods
2.1 Prostate Cancer and SPECT
SPECT imaging of the prostate is dominated by 111In-labeled monoclonal antibodies
and 99mTc-labeled peptides. The most widely used compound is ProstaScintTM, an 111Inlabeled murine monoclonal antibody directed against prostate specific membrane antigen.
The acquisition and interpretation of the ProstaScintTM images are technically demanding
due to radiopharmaceutical uptake by background organs—it takes several days to
accumulate in “target” tissue and clear from “background” tissue. Nonetheless, the
reported sensitivity and specificity of ProstaScintTM imaging for detection of pelvic
lymph node metastases is 62% and 72%, respectively [7].
2.2 Transrectal Single Photon Camera
The design concept that we explored for a single photon camera optimized for
imaging is a transrectal camera. The imaging performance of a single photon camera is
dominated by its collimator. As the parallel-hole collimator efficiency is independent of
the collimator to source distance but the spatial resolution is linearly proportional to the
collimator to source distance, optimal performance is achieved by placing the camera as
close to the prostate as possible. This suggests a transrectal geometry, as shown in
Figure 1, as it places the front face of the collimator less than a centimeter from the
prostate.
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Figure 1. Geometry for transrectal imaging of the prostate with a single gamma camera.
The camera is inserted transrectally, bringing the front face of the camera
close (within ~1 cm) to the prostate. The lines running through the prostate in
this figure indicate the direction that is “viewed” by the camera. The count
rate in each pixel in the camera is proportional to the integrated activity along
one of the lines.
A suitable camera geometry is shown in Figure 2. The gamma ray detector is an 8x8
array of pixels, each 3 mm square. The detector could either be a solid state detector
material (such as cadmium zinc telluride [8-10]) or a scintillator / photodetector
combination (such as CsI:Tl coupled to a PIN photodiode array [11, 12]). The detector
array would sit behind a collimator. For optimal spatial resolution, the collimator would
have square holes that are matched to each element in the detector array. To minimize the
septal thickness and ease manufacturing, the collimator would be made of thin sheets of
tungsten with holes etched via photolithography, then stacked and glued together. As the
number of individual detector elements is relatively high and the leads between the
detector elements and their amplifiers must be short (to minimize electronic noise), the
electronics must be part of the camera head, and would take the form of a custom
integrated circuit mounted on a printed circuit board. This imaging head would be
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mounted on a support structure that would provide mechanical support for insertion into
the rectum and electrical connectivity.
Figure 2. Geometry of a transrectal single gamma camera. The active components of
camera are a gamma imager, readout electronics, and a collimator. They are
mounted on a mechanical support structure and are encased in shielding (not
shown) that blocks gamma rays that do not pass through the collimator.
However, this design has a number of drawbacks. The size of the camera is likely to
make it uncomfortable for the patient and so poorly tolerated. Since the camera is likely
to be imaging 111In-ProstaScintTM, whose emissions are 171 keV and 245 keV, the
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desired detector thickness is 1.5–2.0 cm in order to have high detection efficiency (the
attenuation length in CsI is 0.9 cm at 245 keV and 0.4 cm at 171 keV). One could
probably consider the minimum thickness to be 0.5 cm. For the collimator, collimators
for imaging 111In typically have an aspect ratio (length : hole diameter) of ~10:1. Given a
3 mm hole diameter, this implies a length of 3 cm, which is rather long. The same aspect
ratio can be obtained without having any collimator holes expose more than one detector
pixel by dividing the 3 mm detector pitch by an integer (i.e., have square holes on a
1.5 mm or 1 mm pitch, which divide the pitch by a factor of 2 and 3 respectively).
However, the septa (thickness between holes) must be ~0.5 mm (assuming a tungsten
collimator), so these reduction factors (smaller holes) are problematic because the
collimator septa either begin to occlude a significant portion of the hole or they become
so thin that the gamma rays can penetrate the septa easily. Thus, the desired collimator
thickness is 3 cm, with 1 cm probably representing the minimum thickness. Finally,
approximately 25 attenuation lengths of shielding must be placed around the entire
camera in order to block gamma rays from striking the detector directly (i.e., without
going through the collimator). At 245 keV, the attenuation length of lead and tungsten are
1.5 mm and 1.1 mm, implying desired shielding thicknesses of 3.75 cm and 2.75 cm
respectively. The minimum conceivable shielding thickness is about 1.5 cm centimeter of
tungsten. Thus, the desired thickness is 6.25 cm and the minimum conceivable thickness
is 3.0 cm.
While there are serious concerns as to whether a 3.0–6.25 cm thick device can be
tolerated by the patient, the most serious drawback is that fact that this type of device is
likely to produce a 2-D planar image instead of a 3-D volumetric image. This is
illustrated in Figure 3, which shows how data for a 3-D image is collected with an
external gamma camera. In order to create a 3-D volumetric image using computed
tomography, you must take a large number (~100) of individual planar projections of the
object, with each projection taken from a different angle. Figure 3 shows this data being
collected at three different projection angles—vertical, 45°, and horizontal. While this is
easily accomplished with an external camera, it is obviously impossible to move a camera
with transrectal geometry around the prostate. It is conceivable that with advanced
collimator designs and processing algorithms, some depth information might be obtained
with a stationary camera [13]. An alternate solution might be to use an external camera,
but such a camera would be little different than a conventional SPECT camera, and thus
not have any advantage for prostate imaging. In short, the transrectal geometry is likely
only to be capable of generating a planar image of the prostate, which will have
significantly less contrast and diagnostic utility than a 3-D volumetric image.
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Figure 3. Projection views necessary for tomography. In order to reconstruct tomographic
(3-dimensional) images, projection data of the object must be taken from a
large number of “views.” Figures a), b), and c) show an external gamma
camera positioned to acquire data from three different views (vertical, 45°,
and horizontal respectively), as well as a projection of the object taken at each
view.
3 PET Methods
3.1 Prostate Cancer and PET
In general, PET has higher sensitivity, spatial resolution, and cost than SPECT. The
most commonly used PET radiopharmaceutical in oncologic imaging is 18Ffluorodeoxyglucose (FDG), and the first PET studies of prostate cancer were performed
with FDG. Unfortunately, FDG is not very prostate specific and bladder accumulation of
radioactivity often obscures prostate tumors [14]. Thus, SPECT imaging with
ProstaScintTM is regarded as being superior to PET imaging with FDG for predicting the
presence of prostate cancer [15], particularly for less aggressive disease [16].
Fortunately, newly developed PET radiopharmaceuticals, such as 11C-methionine,
11
C-choline, 11C-acetate, or 18F-fluorocholine, have recently demonstrated outstanding
results in the sensitive detection of prostate cancer. Hara and co-workers find that: 11Ccholine clears the blood quickly; its uptake in prostate tumors provides excellent
tumor/normal contrast; and bladder accumulation is minimal [17]. Therefore, 11C-choline
is an extremely attractive PET tracer for imaging prostate tumors [18-27]. Figure 4 shows
11
C-choline images of a prostate cancer patient before and after therapy, demonstrating
the ability to detect prostate carcinoma and follow therapy efficacy using 11C-choline.
Several other radiopharmaceuticals are also currently under investigation for prostate
cancer imaging, including 11C-acetate [28-30], 11C-methionine [31-34], and 18F-
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fluorocholine [35, 36]. Therefore, we were also motivated to also consider PET camera
designs that are optimized for imaging the prostate.
Figure 4. PET images of choline cancer. 11C-choline image of a patient with prostate
cancer before (a) and after (b) treatment. These grayscale images indicate a high (white)
uptake in the prostate center compared to a low (gray) uptake elsewhere. Images provided
by Hara and co-workers [17].
3.2 Transrectal PET Camera
The first prostate-specific PET camera design that we considered is shown in
Figure 5. Like the single photon camera, this design employs a transrectal probe, but
because PET requires detection of a pair of annihilation photons, it also includes an
external bank of detectors. The advantages of this geometry over conventional PET
geometries parallel those for the transrectal single gamma system. By placing a detector
transrectally, it is extremely close to the prostate, and the external bank of detectors is
also placed as close to the patient’s abdomen as possible. This increases the sensitivity
and also has the potential to improve the spatial resolution. When the object is much
closer to one detector in a PET camera than the other, the spatial resolution is determined
by the closest detector. If this detector has higher resolution than conventional detectors
(while maintaining high efficiency), then high-resolution, high-efficiency imaging can be
achieved [37]. As the volume of the transrectal detector is much smaller than the volume
of the external detector bank, one can use a transrectal detector that costs considerably
more or is considerably more complex than a conventional module without affecting the
commercial viability of the design. In fact, this design uses a factor of ~3 less detector
volume than a conventional PET camera.
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Figure 5. Geometry for transrectal imaging of the prostate with a PET camera. One PET
detector module is inserted transrectally, bringing the front face of that
module close (within ~1 cm) to the prostate. The other end of each chord is
defined by an external bank of PET detector modules The lines running
through the prostate in this figure indicate the directions that are “viewed” by
the camera. The count rate in each chord in the camera is proportional to the
integrated activity along one of the lines.
The size issues discussed with the transrectal single photon imager also exist for the
transrectal PET probe, but are less of a problem. This is because PET relies on electrical
collimation rather than mechanical collimation, so there is no need for a collimator and
the shielding requirements are much reduced. However, the requirements for this PET
probe are significantly higher than they are for conventional PET detector modules. First,
it will be exposed to a much higher gamma flux than conventional PET detector modules,
as it is much closer to the patient than detectors in a conventional PET camera and there
is substantially less shielding. Therefore, the transrectal PET probe must have much
higher rate capability (and therefore much lower dead time) than conventional PET
detector modules. In addition, the probe must be a “volumetric” detector. That is, most
PET detectors identify the position of each 511 keV photon interaction accurately (within
~5 mm) in two coordinates (in a cylindrical camera geometry, the axial and azimuthal
directions), but significantly less accurately (~25 mm) in the third coordinate. Thus, it is
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essentially a 2-D detector. In order to maintain high spatial resolution, the transrectal
probe needs to measure the third coordinate with accuracy similar to what is measured in
the other two coordinates. This can be solved by using a detector that measures the
“depth-of-interaction,” and a number of detector designs have been proposed [38-46] that
could potentially provide the requisite performance.
However, there are several drawbacks to this design. Building a detector probe that
can simultaneously achieve high resolution, high efficiency, and meet the count rate /
dead time requirements is extremely challenging. There are also some potential patient
safety issues. Many of the potentially suitable detector designs utilize photomultiplier
tubes, which have fragile glass envelopes and require operating voltages of
approximately one kilovolt. The design must be robust enough that no harm will come to
the patient even if the photomultiplier tube were to break while inserted into the patient’s
rectum. However, the largest drawback of this design is the same as the largest drawback
in the transrectal single photon probe design, which is the type of image produced. This
design can only collect data from a single (or small number of) projection angles around
the prostate, and so will produce images that are little more than two dimensional.
Although the radiotracer contrast in the patient when using PET radiopharmaceuticals is
better than 111In-ProstaScintTM, the resulting images are unlikely to be of diagnostic
quality.
3.3 External PET Camera
The final prostate-specific PET camera design that we considered is sketched in
Figure 6. It consists of two external banks of detector modules arranged in an elliptical
geometry (45 cm minor axis, 70 cm major axis). This geometry brings the detector
modules as close to the prostate as possible, which increases the sensitivity. The axial
extent is only 8 cm, which is approximately half that of a conventional PET camera. This
provides better shielding, which reduces the number of scatter and random background
events. This design requires about one-fourth of the detector volume of a conventional
PET camera, which significantly reduces the cost.
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Figure 6. Geometry for the prostate-specific PET camera. (a) Drawing of a transaxial
view through prostate, showing the patient centered between two PET
detector banks. The individual detector modules are angled to point towards
the prostate. (b) Drawing of the sagittal view. The bottom arc is fixed below
the patient bed, whereas the top arc adjusts vertically for patient access and
compactness. Both detector banks are tilted and positioned as close as possible
to the prostate, which improves sensitivity and minimizes attenuation.
The bottom detector bank is fixed below the patient bed, whereas the top bank can
be moved upward for patient access and downward for maximum sensitivity. Each bank
consists of 2 axial rows of 20 Siemens/CTI EXACT HR+ block detectors for a total of 80
detectors per camera. We use these HR+ block detectors because they are three
attenuation lengths thick for good detection efficiency with narrow detector elements for
good spatial resolution. The detector modules point toward the center of the camera (i.e.,
close to where the prostate is located) to minimize penetration artifacts. However, this
geometry accentuates the penetration artifacts for sources located near the edge of the
camera field of view. In short, we sacrifice the size of the volume that is imaged with
good spatial resolution to a cylinder at the center of the camera that is approximately
15 cm in diameter and 8 cm high, and in return obtain better sensitivity, lower scatter and
randoms backgrounds, and lower cost. A patient of average size is not fully encircled by
detector modules, which results in irregular and incomplete sampling due to the side
gaps. However, the angular coverage is nearly complete (unlike the transrectal designs),
so we are able to create 3-D volumetric images. Using a 3D iterative penalized maximum
likelihood reconstruction algorithm with randoms and attenuation correction, we are able
to reconstruct nearly artifact free images in the region of interest [47-50].
As this design concept seemed the most advantageous to us, we constructed a
camera based on this concept and have described its performance in detail [48, 50]. We
summarize the main performance measurements here. The sensitivity of a point source in
the center is 946 cps/Ci (2.6%) in 3D mode, which is equivalent to the sensitivity of the
EXACT HR in 3D mode. Figure 7a shows a reconstructed image of a 37-line source
phantom. We are able to resolve line sources separated by 5 mm when placed at a
diameter of 8 cm and 16 cm, demonstrating good spatial resolution in a central region
easily large enough to image the prostate and prostate bed. The spatial resolution,
measured by imaging a 20-gauge needle filled with 18F, is 4 mm full width at half
maximum (fwhm) near the center and increases only to 6 mm fwhm for line sources
placed at a 20 cm from the center. Figure 7b shows a modified NEMA body phantom.
Six spheres are placed on a 6 cm radius in the transaxial plane, surrounding a central
sphere. The two largest spheres in the outer ring were filled with non-radioactive water.
The remaining five spheres were filled with 18F solution with an initial activity density of
1.1 Ci/ml, which was nine times higher than the torso background. Figure 7c shows the
reconstructed images (central slices) of the phantom. All seven spheres can be resolved.
Some artifacts are seen due to incomplete sampling, and we are working to reduce these
artifacts.
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Figure 7. Images taken with the prostate-specific PET camera. (a) Reconstructed image
of 37-line source phantom (each source has a 5 cm axial extent). In the transverse plane,
the single line sources are 2, 4, 6, and 8 cm from the central line. Clusters of four line
sources are placed 4 and 8 cm radially from the central line. The four line sources in each
cluster are spaced 8, 6, 5, and 4 mm apart (clockwise from the 8 mm labeled clusters).
The phantom was filled with 18F at an initial activity of 0.8 mCi, imaged for 2 hours, and
4.4x107 counts collected. The phantom was centered in the camera. Voxel size equals 2
mm x 2 mm x 2 mm. (b) Modified NEMA body phantom. Six spheres are placed on a 6
cm radius in the transaxial plane (37, 28, 22, 17, 13, and 10 mm diameters), and a 28 mm
diameter sphere is placed in the center. All seven spheres have a common axial center
line. (c) Reconstructed images (central slices) of the modified body phantom, which was
centered in the PET camera. Initial 18F activity density was 1.1 Ci/ml in five spheres
(shown in white) and 0.12 Ci/ml in background torso.
4 Conclusions
In recent years both SPECT and PET radiopharmaceuticals have been developed that
are designed for functional imaging of prostate cancer. These developments have led us
to explore nuclear medical imaging camera systems that are optimized for imaging the
prostate. The overall philosophy is that by optimizing for imaging a single disease /
organ, superior tradeoffs in the efficiency, spatial resolution, background levels, and cost
can be obtained. Because of the extremely close proximity to the prostate, designs (both
PET and SPECT) involving transrectal probes promise the best combination of spatial
resolution and efficiency. However, the angular sampling in these designs is quite
limited, making it likely that only planar (as opposed to volumetric) images could be
obtained with them. In addition, it is nontrivial (although not impossible) to make
transrectal probes that meet the performance requirements and still are small enough to
be well tolerated by the patient. We therefore identified a PET camera designed based on
a pair of external elliptical detector banks as the most promising geometry, as it promised
volumetric images with better background rejection and lower cost (but similar spatial
resolution and efficiency) than conventional multi-purpose PET cameras. We have
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constructed and performed initial characterization tests on such a camera, and find that it
performs as expected.
5 Acknowledgement
We would like to thank Dr. Thomas Budinger, Dr. Stephen Derenzo, Dr. Ronald
Huesman, Dr. Jinyi Qi, and Dr. Jicun Hu of Lawrence Berkeley National Laboratory for
many informative discussions. This work was supported in part by the Director, Office of
Science, Office of Biological and Environmental Research, Medical Science Division,
U.S. Department of Energy under Contract No. DE-AC02-05CH11231, in part by
Department of Defense grant number DAMD17-02-1-0081, and in part by National
Institute for Biomedical Imaging and Bioengineering grant numbers R01-EB-00194 and
R01-HL-071253.
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