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Basic Science Investigations |
Department of Radiology, University of Pennsylvania, Philadelphia, Pennsylvania
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ABSTRACT |
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 TOP ABSTRACT INTRODUCTIONMATERIALS AND METHODSRESULTSDISCUSSIONCONCLUSIONREFERENCES |
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Key Words: PET • performance measurements • C-PET • National Electrical Manufacturers Association • International Electrotechnical Commission • NaI PET
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INTRODUCTION |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODSRESULTSDISCUSSIONCONCLUSIONREFERENCES |
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-rays).
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MATERIALS AND METHODS |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODSRESULTSDISCUSSIONCONCLUSIONREFERENCES |
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An energy correction is applied to compensate for local variations in the measured energy of a scintillation event, resulting in more uniform energy resolution over the detector and a system energy resolution of 11%. Energy discrimination is performed by applying a lower and upper energy threshold of 435 and 665 keV, respectively. As with other position-sensitive detectors, the calculated position does not correspond exactly to the point at which the scintillation occurred. To remove systematic errors, an experimentally determined 2-dimensional position offset is added to the calculated position. This procedure is called distortion removal (9).
At this stage, each event is binned into "4-dimensional" projection coordinates. Optionally, the events can be stored in list mode (event-by-event) format, which is particularly suitable for sparse data and for research purposes. The structure of the sinogram files is typically 128 (radial bins) by 96 (azimuthal angles), sorted into 7 tilt angles (out-of-plane or polar angle 
). For brain imaging, the transverse FOV is 256 mm and slice separation is 2 mm, leading to 128 slices. For whole-body imaging, the transverse FOV is 576 mm and slice separation is 4 mm, leading to 64 slices. For research purposes, the sinogram size can be increased to 256 by 192, the number of tilt angles can be varied from 1 to 15, and the transaxial FOV can be reduced to 128 mm. This allows radial sampling from 0.5 mm per bin to 4.5 mm per bin. The bin width is constant for all radial positions.
Before image reconstruction, scatter and random events are corrected using background subtraction. The background is assumed to have either a flat shape (uniform background) or the shape of a parabola in the radial direction (nonuniform background). The resulting background function is fit to the data outside the body contour in the sinogram and is then subtracted from the sinogram, with negative values being set to zero. Background caused by random coincidences tends to be flat (uniform) in the sinogram, whereas background caused by scattered radiation has a curved profile. By varying the curvature of the fitted parabola, one can adjust the calculated background to approximately match the actual background.
Measured attenuation correction is applied after Fourier rebinning (10,11), which converts the 4-dimensional sinograms into stacked 2-dimensional sinograms. Transmission scanning is performed using a 137Cs single-event source of 662 keV. The point source is axially centered in the FOV, and 1 transmission scan covers an axial FOV of 112 mm. Oblique lines of response are rebinned using the single-slice rebinning method (12). This might cause problems if the object density changes significantly within a small, off-center region and work on possible solutions, such as Fourier rebinning or fanbeam reconstruction, is in progress. The patient bed is moved between rotations for complete coverage of the scanner FOV. Because the source is axially centered, the end positions require additional transmission scans (e.g., a single bed position requires 3 transmission scans, whereas 5 bed positions with a pitch of 112 mm require 7 transmission scans). In this case, the reconstructed axial length is 70.4 cm. Transmission scanning can be performed after injection because the 511-keV events from the patient can be separated from the 662-keV events from the single-event source by energy discrimination. The lower- and upper-level discriminators are set at 595 and 860 keV, respectively, during transmission scanning. In addition, an emission contamination, or mock scan, is acquired to correct for any 511-keV events that are measured in the transmission energy window (13). Optionally, the reconstructed transmission image can be segmented, or remapped (14), to reduce the impact of noise and scatter. Typical scan durations at each bed position are 6 min for an emission scan, 55 s for a transmission scan (for 1 rotation of the 137Cs source), and 22 s for the mock scan. For whole-body imaging, the transmission and emission scans are interleaved as the patient is moved through the gantry. The axial sensitivity profile of a single bed position is triangular, and the overlap between 2 neighboring bed positions is approximately 50% to achieve an axially uniform sensitivity profile.
Image reconstruction for whole-body and brain studies is performed using an iterative algorithm that is based on the ordered-subsets expectation maximization method (15) or the row action maximum-likelihood algorithm (RAMLA) (16,17). The missing data in the gaps between the detectors are estimated using the constrained Fourier technique (18) and are needed when using Fourier rebinning to reduce the data to 2-dimensional sinograms before image reconstruction. The reconstructed volume image has the same number of slices and slice separation as the raw sinogram data and an image pixel size of 2.0 mm (brain) or 4.0 mm (whole body). In combination with a 128-mm transaxial FOV and large sinograms, the smallest achievable image pixel size is 0.5 mm.
Performance Measurements
The task of performance measurements is to define an experimental setup that allows one to determine the imaging characteristics of a scanner, to compare different scanners, and to understand and predict the scanner behavior for patient studies. The measurements described here mainly follow the guidelines outlined in the National Electrical Manufacturers Association (NEMA) NU2-1994 standard on performance measurements of positron emission tomographs (19) and standard 61675-1 of the International Electrotechnical Commission (IEC) (20). Furthermore, we performed some of the measurements recently adopted by NEMA to update the existing standard to accommodate a new generation of 3-dimensional PET scanners with longer axial FOVs (>17 cm). We refer to the updated standard as NEMA NU2-2001.
Spatial Resolution.
The axial resolution was determined by a series of point source (all dimensions < 1 mm) measurements in air. The NU2-2001 standard also uses the point source to determine the transaxial resolution, whereas the IEC and NU2-1994 standards require line source (transaxial dimensions < 1 mm) measurements in air. The point source was positioned in the axial center of the FOV (z = 0 cm) and halfway between the axial center and the edge of the FOV (z = 6.4 cm). Transaxially, the point source was positioned at (x,y) = (0,0), (10,0), and (0,10) cm. The line source (length, approximately 5 cm) was positioned parallel to the axis of the tomograph at radial positions of r = 0, 5, 10, 15, 20, and 25 cm. Because of its continuous crystals, the C-PET scanner has no "sweet spot" at (0,0), unlike PET scanners with discrete crystals that are oversampled in the center. Resolution measured at the position (0,0) is, therefore, equivalent to results obtained with a radial offset of 1 cm as prescribed in the standards. For the central position, a sinogram sampling of 0.5 mm per bin and an image pixel size of 0.5 mm were used. For the off-center position at r = 10 cm, a sampling of 1.0 mm per bin and a pixel size of 1.0 mm were used. In both cases, the axial sampling was 2 mm per slice. The images were reconstructed using filtered backprojection, with a ramp filter and a cutoff at the Nyquist frequency.
In accord with the IEC and NU2-1994 standards, horizontal, vertical, and axial profiles (1 voxel wide) were drawn through the pixel with the maximum value in the image. The full width at half maximum (FWHM) was determined using linear interpolation; the equivalent width, EW, was calculated according to the IEC protocol as EW=
i Ci . pixel size/maximum pixel value, where i Ci is the sum of the counts in the profile between the limits defined by 1/20 of the maximum pixel value on both sides of the peak (20). The results were averaged over the horizontal and vertical directions for source positions at equal radial distances r from the center of the FOV. For the line source, the transaxial resolution was also averaged over 16 slices.
Scatter Fraction.
The NU2-1994 and IEC scatter fraction is estimated using a single 18F line source, which is scanned in a water-filled cylinder phantom (diameter [ø] = 20 cm, length = 19 cm) at 3 radial positions (r = 0, 45, and 90 mm). The sinogram data were analyzed according to NU2-1994 (19) and IEC (20), whereby the IEC standards typically lead to a higher scatter fraction (21). Because the C-PET scanner has an axial FOV of 25.6 cm, the test phantom (19 cm in interior length) does not cover the whole axial FOV and values for only the central 17 cm of the axial FOV can be determined (19,20). Also, the measurement does not account for scatter contamination from activity outside the FOV that is present in patient studies.
To overcome the weakness of the previous standards, we investigated a new phantom (NU2-2001) that consists of a 70-cm line source in a polyethylene cylinder (ø = 20 cm) with a water equivalent density of 1.0 g/mL. For ease of handling, the cylinder is cut into 4 segments. The tube containing the activity has a length of 70 cm and a volume of approximately 5.7 mL. The volume of the phantom body is 22 L. The radial position r = 45 mm for the line source is assumed to have scatter representative of that in a homogeneous activity distribution in a phantom of the same size. In the NU2-2001 protocol, the sinogram data are analyzed similarly to the NU2-1994 method.
Sensitivity.
The sensitivity expresses the correlation between activity within the FOV and the number of acquired counts in the absence of dead-time effects. The NU2-1994 and IEC protocols use a 19-cm-long cylinder (ø = 20 cm) for this measurement. There are several concerns about this method. The counts are attenuated within the phantom, and measured counts contain scattered events that are then corrected by multiplying the individual slice sensitivities Si by (1 - SFi), where SFi is the relative scatter fraction for the slice i. Depending on the standard used (NU2-1994 or IEC), this leads to different values for the sensitivity; in addition, the analysis is limited to the central 17 cm of the axial FOV. Because the C-PET scanner has an axial FOV of 25.6 cm, this method will lead to an underestimation of the actual sensitivity. Therefore, we also measured the absolute sensitivity using a phantom similar to that proposed by Bailey et al. (22) (NU2-2001). A 70-cm-long steel tube (inner ø = 3.9 mm, outer ø = 6.4 mm) and 4 concentric sleeves (wall thickness, 1.25 mm) were used. The gap between the sleeves was 0.35 mm. The sensitivity was measured for different wall thicknesses (5 measurements) and extrapolated to zero-wall thickness. The advantages of this method are no significant self-attenuation or scatter contamination, simplified measurement from a physical point of view, and suitability for scanners with a long axial FOV.
Count-Rate Behavior.
The count-rate measurement indicates the relationship between acquired counts and activity level. The count-rate behavior was measured with 5 different phantoms, whereby the phantoms were scanned over several half-lives of the decaying isotope to cover a broad range of counting situations. The first phantom was a 19-cm-long uniform cylinder (ø = 20 cm, volume = 5.8 L, NU2-1994 and IEC). The second and third phantoms were the IEC cardiac and abdomen count-rate phantoms (active volume = 45 mL) (20), whereby the phantom was scanned without arms, because we typically obtain our whole-body scans with arms outside the FOV. Because all 3 phantoms have a length shorter than the axial FOV, their predictive power with regard to clinical scanning is naturally limited. Therefore, we also measured a 70-cm-long uniform cylinder (ø = 20 cm, volume = 22 L). This phantom is long enough to address issues arising from activity outside the FOV, as well. Given the size of this phantom, we had difficulty preparing a solution of uniform activity and handling the phantom. The simplified 70-cm line source phantom from the scatter measurement has also been adopted by NU2-2001 for the count-rate measurement.
Data from the 19-cm-long uniform cylinder were analyzed according to the IEC and NU-2 standards, meaning that only the central 17 cm of the axial FOV and only events within a 12-cm radius were considered. For the cardiac and abdomen phantoms, only the central 17 cm were considered, but no transaxial restrictions were applied. For these 2 measurements, the background subtraction method was used to estimate scattered and random events. Data from the 70-cm uniform cylinder were analyzed using background subtraction, and only events within a 12-cm radius were considered. Analysis of the 70-cm line source data was analogous to the scatter fraction analysis, allowing a simple, potentially more accurate estimate of true, scattered, and random events. For the comparison with patient count-rate data, all phantom data were also processed following a clinical protocol, meaning that background subtraction was used to correct for scatter and random events, and all events within the sinograms were considered (no transaxial restrictions). Therefore, the total count rates are expected to be higher than the count rates found according to the performance standards. To compare the different phantom and clinical data, we plotted the count rates against single count rates, which correlate quasilinearly with the activity seen by the scanner.
Count-Loss Correction.
To quantitatively measure source activity distributions under widely varying count-rate conditions, the PET scanner must compensate for dead-time losses and random events. For this test, the count-rate data described above were reconstructed with all corrections applied, including those for count losses and physical decay. Regions of interest (ROI) were drawn on each image, and the counts were plotted as a function of the activity concentration. Ideally, all ROIs should have the same number of counts, leading to a line parallel to the abscissa. Deviations from this parallel line indicate errors in the count-loss correction or the random-event subtraction.
Recovery Coefficients.
Measurement of the recovery coefficients is required only by the IEC standard. A set of 6 hollow spheres with different diameters (ø = 10, 13, 17, 22, 28, and 37 mm) filled with 18F solutions of identical activity concentrations was placed in a water-filled cylinder (ø = 20 cm, length = 19 cm) (20), which was centered in the transverse FOV. The centers of the spheres were arranged to lie in 1 plane in the center of the phantom. Two axial positions were measured: spheres centered in the axial direction (z = 0 cm) and spheres halfway between the axial center and the edge of the axial FOV (z = 6.4 cm). The recovery coefficients RCi = Ci/C6 were calculated as previously described (21), where Ci were the ROI counts per pixel per second for sphere i = 1,..., 6, with sphere 6 being the largest sphere. Data were acquired with a 4-mm slice spacing and a 256-mm transverse FOV, and images were reconstructed using filtered backprojection.
Image-Quality Phantom.
Existing performance measurements describe mainly a single parameter, often requiring best possible statistics. Under clinical conditions, the scan duration and activity concentrations are constrained, and the count statistics are limited. A measurement has been designed by NU2-2001 to assess the trade-off between different parameters (e.g., sensitivity, scan duration, activity level, shielding, and septa) (23). The task is to scan a 1-m axial length within 1 h, including transmission and emission scans and using clinically realistic activity levels. The phantom is the IEC whole-body phantom with the sphere inserts of the IEC recovery coefficient phantom and a cylindric foam insert (ø = 50 mm; P = 0.3 g/cm3). Because the IEC whole-body phantom is only 20 cm long and activity outside the FOV is expected to degrade image quality, the phantom is combined with the 70-cm-long line source count-rate phantom. A sketch of the phantom configuration is given in Figure 2. The image-quality phantom test supersedes the NU2-1994 tests of scatter correction, attenuation correction, and image uniformity.
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ROIs with diameters equal to the physical inner diameter of each sphere were drawn on the slice through the centers of the sphere. Twelve ROIs of the same sizes as those for the spheres were drawn throughout the background in the central slice and in the slices ±0.8 and ±2.0 cm away. The coefficient of variation (covariance) of these 60 background ROIs was determined for each sphere size as a measure of the background variability (CRCbkgd). The hot-sphere contrast recovery coefficient (CRChot) was calculated as:
 | (Eq. 1) |
For the cold spheres and the foam insert, the contrast recovery coefficient (CRCcold) was calculated as:
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Clinical Images.
To show the quality of clinical images, we selected a representative 18F-FDG whole-body study, which constitutes approximately 85% of our studies.
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RESULTS |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODSRESULTSDISCUSSIONCONCLUSIONREFERENCES |
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Scatter Fraction
The average scatter fraction according to NU2-1994 is 25.0%; according to IEC it is 31.6%, and for the 70-cm line source it is 35.0%. Because of contributions from out-of-FOV activity, the 70-cm phantom has the highest scatter fraction. The differences in the data analysis procedures for the IEC and NU2-1994 standards cause a significant difference in the calculated scatter fraction, whereby the IEC value is 26% higher than the NU2-1994 value. The NU2-2001 slice scatter fractions, given in Figure 3, have a maximum in the center of the axial FOV, from which they slowly decrease toward the edge of the FOV. As expected, the scatter-fraction curve is symmetric.
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Count-Rate Behavior
The count-rate behavior for the 5 different phantoms is shown in Figures 4A–4E. The plots show that the count-rate curves differ significantly for the different phantoms. Figure 4A shows also that, because of the different scatter fractions measured according to IEC or NU2-1994, the maximum true count rate and noise equivalent count rate (NEC) are higher by 9.6% and 20.0%, respectively, for the NU2-1994 measurement than for the IEC measurement. The total and the random count rates are not affected by the difference in scatter fractions. The two 70-cm-long phantoms (Figs. 4B and 4C) are the only 2 phantoms that have a similar count-rate behavior. Figure 4F compares phantom count rates with actual patient count rates. The count rates of the 70-cm-long phantoms correspond to those for patient data (cardiac and whole-body scans), whereas the 19-cm-long cylinder seems to give an upper limit for brain and head scans. Because we have a dedicated brain scanner in our PET center, we rarely perform brain imaging on the C-PET scanner, and only a few data points for brain imaging are available. The IEC abdomen phantom gives a count rate between the count rates for the 19- and 70-cm-long phantoms, whereas the IEC cardiac phantom has a count-rate behavior that does not correlate well with observed patient count rates. These values are derived mainly from FDG whole-body scans but include some cardiac studies as well. The discrepancy is caused mainly by the small amount of attenuating material in the IEC cardiac phantom and the unrealistically high ratio of true events to single events compared with patient scans. In addition, because the activity is in a small insert (volume, 45 mL), the activity concentrations in Figures 4D and 4E are much higher than in Figures 4A–4C. Furthermore, the count-rate curves show that the maximum NEC values are constrained on the C-PET scanner more by dead-time effects and less by the random-event count rate. The random events are always less than one third of the true count rate at the peak value of the true count rate.
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DISCUSSION |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODSRESULTSDISCUSSIONCONCLUSIONREFERENCES |
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The measured recovery coefficients are comparable with those of other whole-body systems. However, the comparison of recovery coefficients and CRCs from the image-quality phantom shows that measurement of a small lesion is much more difficult if the background is hot instead of cold and if the phantom is nonuniform. The low CRCs also indicate that the scatter and random-event corrections need to be improved.
Several projects to improve the C-PET scanner are under way. These include a scatter correction using a dual-energy window method (25), which takes advantage of the good energy resolution of NaI(Tl). The hardware has already been implemented, and the software implementation is soon to follow. In combination with measured random-event contamination using a delayed coincidence timing window, which will also be available soon, this method will make the background subtraction obsolete and improve correction of the data. Recent changes to the acquisition hardware and software allow an increased sinogram size (192 x 256) and larger images (256 x 256 or 288 x 288 pixels per slice). These will improve image resolution and should lead to higher measured recovery coefficients. We are investigating the use of 3-dimensional RAMLA (25). Initial results show that images reconstructed with 3-dimensional RAMLA have higher contrast and lower noise, compared with images reconstructed with ordered-subsets expectation maximization (16).
As expected, the differences in the performance standards lead to differences in the results. In particular, the method used to determine scatter in the IEC and NU2-1994 measurements affects not only scatter fraction but also sensitivity and count-rate measurements. Use of phantoms whose axial extent is shorter than the axial FOV of the scanner underestimates the actual count-rate capabilities of the scanner, compared with a scanner with a shorter axial FOV. This is also true for the NU2-1994 and IEC sensitivity values, which are underestimated because of the shortness of the phantom.
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CONCLUSION |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODSRESULTSDISCUSSIONCONCLUSIONREFERENCES |
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Concerning the performance standards, the current NU2-1994 and IEC measurements (which use 19-cm-long phantoms) characterize the best performance for PET imaging, whereas the new proposed measurements (using longer phantoms for scatter, count-rate, and sensitivity measurements) better characterize the performance under clinical conditions and are better suited to characterize 3-dimensional whole-body scanning. Specifically, the clinical count-rate behavior is better predicted by the 70-cm-long phantom than by the IEC whole-body phantoms. The use of 70-cm-long phantoms will also allow better comparison between scanners with different axial FOVs. In addition, measurements and data analysis are simplified in the NU2-2001 standard, which permits faster testing with little loss of information with regard to scanner performance.
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ACKNOWLEDGMENTS |
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FOOTNOTES |
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For correspondence or reprints contact: Lars-Eric Adam, PhD, Department of Radiology, Hospital of the University of Pennsylvania, 3400 Spruce St., Philadelphia, PA 19104.
Robin J. Smith died in April 2000.
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REFERENCES |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODSRESULTSDISCUSSIONCONCLUSIONREFERENCES |
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