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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5255_Библиотеки_им_академика_М_И_Перельмана
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a set of 2D equivalent projections then any of the 2D reconstruction algorithms can be used
for image reconstruction.
Positron emission tomography: physical principles, instrumentation and performance evaluation
PET/CT
The introduction of combined PET/CT scanners in a single gantry has made a paradigm
shift in the clinical application of this imaging modality (28). The co-registration of PET
and CT data enables regions of increased radioactive uptake on the PET image to be directly
correlated with their anatomic locations on the CT scan, thereby improving the sensitivity
and specificity of PET for lesion detection and characterization (36-38). The impact of this
technology is dramatic such that within 6 years of the introduction of the first PET/CT
scanners, they are accounting for more than 90% of all PET scanner sales in the market
worldwide (38). With attenuation correction performed by the CT component, PET/CT can
provide better quality images over shorter examination times than conventional PET.
Advantages of combined PET/CT over PET and CT acquired on separate devices include
more accurate CT and PET data co-registration, improved lesion localization, consolidation
of patients’ studies, and reduced scan times. In addition, the PET/CT gantry opening is large
enough to accommodate a flat pallet for imaging patients for the purpose of radiotherapy
treatment planning. It may potentially offer more accurate assessment of tumor volume
influencing the therapeutic management of patients by radiotherapy (2,39-41).
Latest versions of PET/CT are providing increased axial field of view (AFOV) with
more detectors that considerably increase the sensitivity and reduce the scan time. As
mentioned before systems are coming up with more detector elements (>30,000) arranged in
multiple detector rings with large AFOV (>20 cm). Efforts are still in process to further
increase the sensitivity by increasing the AFOV thereby reducing the scan time or activity
dose to the patient or both. It is now claimed that a patient can be imaged in the latest PET
system within 10 minutes. On the CT imaging part, today 64 slice CT is available though for
PET/CT earlier versions of CT may be sufficient for co-registration and attenuation correction
(42). The PET/CT technology has found widespread clinical applications but still some
technical challenges remain to be addressed including the physiologic motion during
acquisitions and the effect of CT contrast agents when using the CT data sets for PET
attenuation correction.
Preclinical PET
PET has been used for research studies in animals for long time now. Most of the applications
were in large-animal models for cardiovascular research and the non-human primate for
neuroscience research. Most of these studies were performed on human PET scanners, and
little work was done in rodent models because the spatial resolution of clinical scanners has
been insufficient for all but the crudest of studies (28). The combination of large PET
scanners and expensive animal models also restricted these types of studies only to the big
research centers. But the development of high-resolution, dedicated animal PET scanners
Positron emission tomography: physical principles, instrumentation and performance evaluation
283
during the 1990s brought PET technology to a whole new set of investigators in the biologic
sciences and, as a pre-clinical tool, to the pharmaceutical and biotechnology industry
(21,24,43-45). Just as with human PET scanners, this technology has progressed from
academic laboratories to small companies and now to the major medical imaging corporations.
Although small animal scanners are based on the same technologies as are used in clinical
systems, the requirements like high resolution, high sensitivity, small field of view and the
imaging environment are quite different, leading to very different tradeoffs in the design
(28). Preclinical PET has now become an established discipline that is extremely useful in
the field of molecular imaging, allowing PET tracers across species, from mouse to patient
(28).
Future Scopes
Time-of-Flight (TOF) PET
In conventional PET, a valid event is formed when the two coincident 511 keV annihilation
photons are detected within some pre-defined timing window, typically on the order of
4.5-10 ns for detectors based on scintillators such as LSO and GSO. The two detectors in
which interactions are measured determine a line along which the original annihilation takes
place. The location of the annihilation site is unknown and can be recovered only by image
reconstruction. The image reconstruction algorithm, with no other information at its disposal,
assumes that all possible locations of the annihilation site on the LOR have equal probability.
In TOF PET, the actual time difference (t) in the arrival of the two annihilation photons at
the respective detectors is recorded (Figure 9). The time difference increases the farther the
annihilation site is from the point midway between the two detectors (x). The utilization of
TOF is known to improve the signal-to-noise ratio in PET images, by reducing the noise
propagation along the LOR during the forward and back-projection steps in image
reconstruction.
Figure 9: Schematic diagram showing the concept of TOF PET system
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Positron emission tomography: physical principles, instrumentation and performance evaluation
The noise reduction translates to an effective sensitivity gain, which has been described
simply as a ratio of D/x, where D is the object diameter, x = c t/2, c is the speed of light
and t is the difference in time of arrival of two photons to respective detectors (timing
resolution).
The first TOF PET systems were built using either CsF or BaF2 scintillators in the early
1980s, though, the concept of TOF was proposed as early as in the 1960s (46,47). These
early systems used a 1-to-1 coupling of the crystal to PMT. The intrinsic timing resolution
with CsF or BaF2 was as good as 350 to 440 ps, however the poor spatial resolution and low
light output hampered their use in TOF PET systems (48). By the early 1990s, whole-body
oncology studies with 18F-FDG became prevalent. Today, there are new scintillation materials
with fast decay timing and high light output that overcome the limitations of CsF and BaF2.
They are LaBr(Ce) and LYSO with similar timing properties as LSO (49,50). Several
groups, including CTI, Donner Laboratory and the University of Pennsylvania (using Philips
LYSO crystal) report the first results of the performance of TOF PET scanners (51-53). The
timing resolution of 600 ps with LYSO crystal can detect an annihilation event within few
centimeters of the location of actual event, which is better than having no information of the
annihilation event along the LOR in non-TOF systems. It appears that the TOF PET scanners
can now be designed to have all the desirable features and high performance of non-TOF
scanners with the added benefit of image improvement. If one wants to use the TOF method
to pinpoint the annihilation site to about 5 mm and completely eliminate the need for image
reconstruction, then the photon arrival times would need to be recorded with a precision of
approximately 30 ps. Detectors and electronics capable of handling such a timing resolution
are presently not available. The current research in PET technology is more focused towards
the development of faster detector materials for better energy resolution and extended axial
field of view for improved sensitivity (54,55). Of late time of flight PET concept is revived
with improved detector technology and it may help in improved scanner design for preclinical
PET with improved sensitivity.
PET/MRI
PET/MRI theoretically provides the same advantages that PET/CT provided to dedicated
whole body PET imaging such as anatomical landmarks, shorter scan durations, and
attenuation correction. In addition, PET/MRI potentially can augment PET imaging with a
wealth of other information such as functional and spectroscopic data that may be helpful in
improving patient management as well as understanding tumor biology. Furthermore, with
PET/MRI this information is obtained at a reduced patient radiation exposure compared to
PET/CT. So in essence, pairing PET with MR can only provide an added advantage over
PET/CT (56). The simultaneous whole body PET/MRI imaging appears to be feasible in the
near future (57).
Positron emission tomography: physical principles, instrumentation and performance evaluation
285
Evaluation of PET performance parameters
Most of the currently available PET scanners have been designed by integrating the dedicated
PET and CT scanners as a dual imaging system (PET/CT). To ensure the quality of scans on
PET/CT scanners, as well as to determine optimal system operation parameters, it is important
to be able to characterize their performance in a reproducible and reliable manner according
to accepted measurement standards. The National Electrical Manufacturers Association
(NEMA) recommended certain measurement standards for PET performance in NU 2-1994
standard, which was updated in 2001 as NU 2-2001. The updated document includes revised
measurements for spatial resolution, scatter fraction, sensitivity, counting rate performance,
and the accuracy of count loss and randoms corrections (58). The NU 2-2001 document also
includes a specification for a new image quality measurement that was not considered in the
original NU 2-1994 standard (59). Because of the prevailing interest in clinical whole-body
scanning, this measurement is intended to simulate realistic clinical conditions for such
scans.
The performance tests for a PET scanner are normally divided into two groups (59). For
measurement in both the groups 18F radionuclide is recommended. In the first group basic
intrinsic measurements of spatial resolution, sensitivity, scatter fraction, count losses
(including noise equivalent count rate) and random coincidences are measured. The second
group includes measurements of the accuracy of corrections for physical effects such as
count losses and random coincidences, and overall image quality.
Measurement of intrinsic performance parameters
Spatial Resolution
The spatial resolution represents the ability of a system to distinguish between two radioactive
point sources in an image. The purpose of the measurement of spatial resolution is to
characterize the widths of the point spread function (PSF) in the reconstructed image of
radioactive sources. The width of the PSF is reported as the full width at half maximum
(FWHM) and full width at tenth maximum (FWTM). The spatial resolution is measured in
the transverse slice in two directions, radially and tangentially, and in the axial direction.
For all scanners, point sources of 18F are imaged in air. The point source consists of a
small quantity of concentrated radioactivity inside a glass capillary with inner diameter less
than 1mm in all three directions. The resolution is measured with the sources at 6 locations.
Two axial positions are selected namely, the center of the axial field of view (AFOV) and a
position one fourth of the axial FOV from the center. For both axial locations, the source is
imaged at 3 positions, (a) x = 0 and y = 1 cm, (b) x = 0 and y = 10 cm, and (c) x = 10 and
y = 0 cm. The data are reconstructed with filtered backprojection using an unapodized filter;
the image pixel size should be smaller than one third of the expected FWHM. No smoothing
of the data is performed. The FWHM and FWTM of the PSFs are determined in all 3
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Positron emission tomography: physical principles, instrumentation and performance evaluation
directions by forming 1-dimensional (1D) response functions through the peak of the
distribution in the 3 orthogonal directions. The width of these profiles at right angles to the
direction of measurement is approximately twice the FWHM in those directions, rather than
a single pixel, to reduce measurement variability. The FWHM and FWTM are calculated by
linear interpolation between adjacent pixels at one half or one tenth of an estimate of the
maximum value of the response function. This maximum value is determined by a parabolic
fit using the peak value and it’s two nearest neighboring points. The radial, tangential, and
axial resolutions for each measurement radius, averaged over both axial positions, are reported.
Sensitivity
The sensitivity of a scanner represents its ability to detect annihilation radiation. PET system
sensitivity, for a point source at the center of the FOV, ranges from 0.2 to 0.5% for 2D
scanners to 2 to 10% for 3D scanners. Whereas the SPECT system sensitivity, is typically 10
and 100 fold lower than 2D and 3D PET scanners, respectively (27). The PET scanners are
usually sensitivity limited resulting in low SNR produced in the reconstructed images. In the
NU 2-2001 standard, the absolute sensitivity of a scanner is measured. Because the emitted
positrons annihilate with electrons to create a pair of -rays, a significant amount of material
must surround the source to ensure annihilation. This surrounding material also attenuates
the created -rays, prohibiting a measurement without interfering attenuation. To arrive at an
attenuation-free value of the sensitivity, successive measurements are made with a uniform
line source surrounded by known absorbers. The sensitivity with no absorber can be
extrapolated from these measurements.
In the NU 2-1994 standard, the sensitivity was measured using a standard source and
reported as the rate of coincidence events for a given activity concentration in that phantom
(20 cm diameter, 19 cm long). For scanners with an axial FOV longer than 19 cm, however,
this measurement could under represent the efficiency of the scanner. For other measurements,
a 70 cm long phantom is used. A 70 cm long uniform cylinder, however, would be impractical
to fill and to handle.
Uniformity
The 3D 68Ge phantom is convenient to check the uniformity of PET detector system. However,
a cylindrical phantom filled with uniform source of 18F is more appropriate for such a
measurement. Adequate counts as per protocol have to be acquired. Non-uniformity is
estimated as per NEMA NU-2, 1994 or more appropriately NEMA NU-2, 2001 protocol
(58-60).
Intrinsic Scatter Fraction
The scattering of annihilation photons leads to falsely positioned coincidence events.
Variations in design cause PET scanners to have different sensitivities to scattered radiation.
Positron emission tomography: physical principles, instrumentation and performance evaluation
287
The intrinsic scatter fraction is a measure of the relative system sensitivity to scatter. For a
given source distribution, a lower scatter fraction is more desirable, regardless of the accuracy
of the method for scatter correction, because correction techniques cannot compensate for
the noise introduced by the unwanted events and can, potentially, add bias to the image. The
scatter fraction is defined as the ratio of scattered events to total events, which are measured
at a sufficiently low counting rate that random coincidences, dead-time effects, and pulse
pileup are negligible. Total events, therefore, are the sum of unscattered events (trues) and
scattered events.
The phantom in this measurement, as well as that of count losses and randoms, comprises
a 20 cm diameter solid polyethylene cylinder with an overall length of 70 cm. The data for
the measurement of intrinsic scatter fraction are taken from the low activity scans of the
measurement of counting rate performance, when the count loss and randoms rates are both
<1% of the true rate. Activity is placed in a line source that is threaded through a hole in the
cylinder at a radius of 4.5 cm and parallel to the central axis. It should be noted that the line
source is sufficiently large to preclude problems with bubbles, despite its length, provided
reasonable care is taken during filling. For consistency, the phantom is rotated such that the
line source is at the lowest position, because the measured result will depend on the relative
orientation of the line source and the bed.
Counting Rate Performance
Most patient PET studies are not performed under conditions of low counting rate losses or
negligible randoms rates. At higher activity levels, coincidence events are lost because of
system dead time, whereas the rate of random coincidences rises. It is necessary to measure
the counting rate performance (both dead-time losses and randoms) as a function of activity
to understand the scanner’s behavior for a wide range of scanning conditions. The 70 cm
polyethylene cylinder with a line source described for the measurement of the intrinsic
scatter fraction is used in this measurement. For this procedure, however, the line source is
filled with a known initial amount of activity sufficiently high that both the peak trues rate
and the peak noise equivalent count rate (NECR) can be measured. This starting activity is
determined empirically for each scanner and acquisition mode; the manufacturer will generally
provide a recommended initial activity. Data are taken until the randoms and dead-time
losses are negligible. To avoid potential problems with nonuniform dead time associated
with the off-center line source completely, tomographic data are acquired. Sinograms are
generated, and oblique sinograms are collapsed into a single sinogram for each slice while
preserving counts by using single-slice rebinning.
Accuracy of Corrections for Count Losses and Randoms
To achieve quantitative measurements of source activity distributions under widely varying
conditions, PET scanners must have the capability to compensate for dead-time losses and
randoms. The accuracy of these corrections, particularly at the highest counting rates
288
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Positron emission tomography: physical principles, instrumentation and performance evaluation
encountered in clinical imaging, is reflected by the bias with which the scanner counts. The
accuracy of corrections for dead-time losses and randoms is determined from the counting
rate performance data. The data are reconstructed with all count-rate dependent corrections
(i.e., for dead-time losses and randoms) applied. A large (18 cm diameter) region of interest
(ROI) is defined, centered on the phantom, and the residual error R as a function of effective
activity concentration is given by:
R
R=
trues
R
extrap
(4)
Where R
is determined from the low counting rate data, where there are negligible
extrap
dead-time losses or randoms, and is assumed to be a linear function of activity. A large ROI
is used, as opposed to a small region centered on the line source, to reflect any inaccuracies
in randoms subtraction over the entire phantom. The errors are calculated for each slice, the
largest and smallest errors over all slices are plotted as a function of effective activity
concentration. The maximum absolute value of the bias is also reported for activity values at
or below the activity level of the peak NECR. For estimation of NECR, both trues (T) and
randoms (R) are measured and displayed on the monitor in modern systems. The value of
scatter fraction needs to be determined. A hollow cylindrical phantom with three line sources
at the three radial positions including the center is filled with water. The measurements are
taken as per NEMA protocol for calculating scatter fraction (S) (58-60). The NECR is
calculated using equation-3.
Image Quality Measurement
Because of the complex interplay of different aspects of system performance, it is desirable
to be able to compare the image quality of different imaging systems for a standardized
imaging situation that simulates a clinical imaging condition. The purpose of this measurement
is to produce images simulating those obtained in a whole-body study with both hot and
cold lesions. Spheres of different diameters are imaged in a simulated body phantom with
nonuniform attenuation. Activity is also present outside the scanner. The phantom consists
of a torso phantom containing hot and cold spheres (wall thickness, 1 mm) in a warm
background. The hot spheres have inner diameters of 1.0, 1.3, 1.7, and 2.2 cm; the cold
spheres have inner diameters of 2.8 and 3.7 cm. A 5 cm diameter insert with an attenuation
coefficient approximately equal to the average value in lung (density, 0.30
g/mL) is also placed in the center of the phantom, and attenuation correction is performed.
The background is filled with 18F at an activity concentration typical of what is seen in
patient FDG studies (370 MBq/70 kg patient). The hot spheres are sequentially filled with
activity concentrations of 8 and 4 times that of the background. This process is efficiently
accomplished by first putting the background activity in one eighth of the background’s
volume, withdrawing sufficient volume to fill the spheres, and then filling the rest of the
Positron emission tomography: physical principles, instrumentation and performance evaluation
289
background with non-radioactive water. After one half-life has elapsed, another sample of
background activity is used to increase the background activity back to the starting level and
reduce the hot sphere activity ratio to 4:1. The line source of the 70 cm long phantom is
filled with sufficient activity to yield an effective activity concentration equal to that of the
background in the torso phantom; the 70 cm long phantom abuts the body phantom to
approximate the clinical situation of having activity outside the scanner FOV.
PET normalization
In PET scanners the detectors may exhibit some non-uniformity of response. The modern
scanners consists of 10,000 to 30,000 detector elements, slight variations among the detector
elements whether in thickness, light emission properties or in the electronics performance
may result in some change in count rates for the same activity. The normalization scan is
performed using a uniform positron-emitting 3D phantom source of 68Ge placed over the
entire axial FOV. This exposes the detector pairs to a uniform photon flux. The 3D phantom
is scanned and a data in the form of a sinogram are acquired to generate the correction
matrix (Figure 10).
Figure 10: A sinogram is shown in fan sum view. The bottom portion of the image on the left side
displays tracking information, which is helpful in diagnosing detector / analogue board problem.
Problem in a particular detector of a block and bucket can be identified easily. The axial slice of
the phantom is shown on the right side showing uniformity across the plane.
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Positron emission tomography: physical principles, instrumentation and performance evaluation
Normalization corrects for the variations in efficiency in line of responses (LORs) in
each slice of the sinogram. Normalization in 2D mode is performed by direct measurement
using a low activity source, for 3D mode indirect measurement is performed using a uniform
phantom. The normalization scans are acquired quarterly, after system maintenance or as
suggested by the manufacturer. During normalization the detected coincidence events are
mapped to specific crystal locations and corrections are made for the crystal drifts. The PET
normalization procedure also balances the gain characteristics of the PMTs in a block,
compensates for PMT gain drift with temperature, adjust for timing delays so events from
all blocks are time-stamped equivalently.
PET/CT co-registration
The PET/CT uses DICOM-based coordinate system for co-registration. Since the image is
acquired in the same gantry without any change in position of the patient, it is easier for the
system to calculate the coordinates of a point in PET and its corresponding values in CT. A
phantom with small hole to accommodate a lead ball can be used to check the accuracy in
co-registration. A small drop of 18F can also be introduced into the hole along with the lead
ball. The CT image provides the coordinates of the lead ball whereas PET provides the
coordinates of the 18F point source. Since both the sources are exactly at identical location
the co-registration can be easily checked.
Figure 11 shows the perfect co-registration on the left and mis-registration on the right
side of the image in Siemens’ Biograph. If either the X or Y calibration values exceed 5 mm,
mechanical re-adjustment of the gantries is required. Gantry offset calibration procedure
must be run to ensure that the field of view (FOV) of the system matches within the PET
and CT gantries.
CT Performance measurements
A CT scanner consists of many electromechanical components. They include specialized xray generators and components designed to localize and define the scan plane to the patient’s
anatomy. Since the x-ray generator and electromechanical components are subject to
miscalibration, mechanical misalignment or other problems, careful testing and performance
measurements are required.
Electromechanical tests
This category includes gantry and patient table motions, scan alignment lights and x-ray
beam collimators, accuracy of table vertical and longitudinal movement, radiation and
sensitivity profile widths and tests on x-ray generator.
Positron emission tomography: physical principles, instrumentation and performance evaluation
Figure 11: Image showing perfect co-registration on the left and mis-registration on the right
side.
Scan localization light accuracy
291
Patient anatomy to be scanned is often defined by scan alignment lights. Alignment lights
may be located within the gantry at the slice plane, outside the gantry at a reference distance
from the scan plane, or both. If both internal and external alignment lights are supplied, and
are independently aligned, both should be tested. Congruence of gantry laser with centre of
imaging plane and gantry tilt accuracy can be verified using ready pack film (61).
Calibrations of table linear scales are verified by moving the table both vertically and
longitudinally in steps of 1 cm using an independent measuring scale. Table indexing accuracy
and reproducibility are tested by irradiating a ready pack film placed perpendicular to the
scan plane, under scanner control longitudinal spacing of 0, 5, 10, 20 and 30 cm using 0.1
cm slice thickness.
Collimation
A CT scan samples a thin slab of tissue; ideally only the imaged slab is irradiated, and its
thickness equals the selected slice width. Inaccuracies arise from scanner limitations and
calibration of collimator settings. Most systems collimate the z axis dimension of the x-ray
beam at both the source (patient collimation) and detectors (post patient collimation). Width
of the imaged slice (from sensitivity profile) is determined by both sets of collimators.
Width of the tissue irradiated (from radiation profile) is a function only of the pre-patient
collimator. Measurements of radiation and sensitivity profiles are best interpreted together.
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