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Special Contribution |
1 Department of Radiology, University of Pennsylvania, Philadelphia, Pennsylvania
2 CTI PET Systems, Knoxville, Tennessee
3 Picker International, Cleveland, Ohio
4 Philips Medical Systems, Milpitas, California
5 Philips Medical Systems, Philadelphia, Pennsylvania
6 Siemens Medical Systems, Hoffman Estates, Illinois
7 General Electric Medical Systems, Milwaukee, Wisconsin
8 Department of Radiology, University of Washington, Seattle, Washington
9 Department of Physics, University of British Columbia, Vancouver, British Columbia, Canada
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ABSTRACT |
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 TOP ABSTRACT INTRODUCTIONMATERIALS AND METHODSRESULTS AND DISCUSSIONCONCLUSIONREFERENCES |
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Key Words: PET • performance measurements
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INTRODUCTION |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODSRESULTS AND DISCUSSIONCONCLUSIONREFERENCES |
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Since the publication of the NU 2-1994 standard, there have been several developments in PET scanner technology. The NU 2-2001 standard is an attempt to respond to this changing technology. In particular, 3-dimensional (3D) imaging has become an important acquisition and reconstruction mode, whereas the previous standard assumed only 2-dimensional (2D) imaging. It became necessary to standardize how oblique lines of response (LORs) would be handled during measurement of scanner performance in 3D mode.
In addition, all coincidence gamma cameras have a very large axial field of view (FOV), typically 30–40 cm. Yet, the standard NEMA phantom defined in NU 2-1994 is only 19-cm long, and the tests using this phantom assumed that the axial FOV of the scanner was 
17 cm. Before 1991 all commercial PET scanners did, in fact, have an axial FOV shorter than 17 cm. Today there are also several dedicated PET systems with an axial FOV as large as 25 cm, so the NU 2-1994 tests cannot adequately accommodate these scanners or the coincidence gamma-camera systems.
An issue related to the axial FOV is that the 20 x 19 cm NEMA phantom was designed to represent a compromise between brain imaging and body imaging. In the last 10 y, whole-body 18F-FDG studies have become the predominant type of PET study performed by most centers, driven partly by the available reimbursement through Medicare and local insurance carriers. Therefore, it is more relevant than before to measure the performance of PET scanners under conditions that better represent whole-body studies. In fact, a brain is much smaller than the NEMA phantom (although the head volume is comparable to the phantom volume), whereas a body is much larger. However, because most PET systems before 1991 operated in 2D mode with septa and with an axial FOV of <17 cm, the counting rate data acquired with the NEMA phantom served as a reasonable test of the scanner performance for both conditions. This is no longer true for a modern 3D PET system without septa, particularly if the axial FOV is >17 cm. Therefore, the most significant change in the NU 2-2001 standard, compared with the NU 2-1994 standard, is the change from a 19-cm-long phantom to a 70-cm-long phantom (albeit with a line source) for several of the tests, including sensitivity, scatter fraction, counting rate performance, and count loss correction. The 70-cm-long phantom is a better approximation to the activity distribution in whole-body studies, because the effects of out-of-field activity (OFA) are included in the performance measurement. The 20 x 19 cm NEMA phantom, on the other hand, continues to be used to test the performance of scanners used primarily for brain imaging, particularly for scanners that are dedicated brain imaging instruments.
Another major change in the updated standard is the addition of an image quality test. This measurement attempts to reproduce a clinical situation (an FDG whole-body study) and to take all aspects of camera performance and clinical imaging into consideration, including image reconstruction. The phantom used for this measurement is the IEC body phantom (3) with hot and cold spheres with an additional cylindric insert to simulate lung, as well as a second phantom to provide activity outside the scanner to simulate body uptake. This test was designed to mimic an FDG tumor study, with typical background activity and hot sphere contrasts encountered in those studies. The test is performed with a standard acquisition protocol that allows comparison between different types of PET scanners.
This article presents a description of each test and discusses the changes from the NU 2-1994 standard, along with illustrative data using several different PET systems. The NEMA Standards Publication NU 2-2001 (4), which provides more detail than presented here, is available from NEMA (1300 N. 17th St., Suite 1847, Rosslyn, VA 22209, www.nema.org). Readers wishing to perform the measurements should consult this document for a complete description of the procedures.
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MATERIALS AND METHODS |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODSRESULTS AND DISCUSSIONCONCLUSIONREFERENCES |
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Spatial Resolution
The spatial resolution of a system represents its ability to distinguish between 2 points of radioactivity 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 compact 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 2 directions, radially and tangentially, and in the axial direction.
For all scanners, point sources of 18F (<1-mm extent in any direction) are imaged in air. The point source consists of a small quantity of concentrated radioactivity inside a glass capillary (wall thickness, <0.5 mm). The resolution is measured with the sources at 6 locations. Two axial positions are selected—namely, the center of the axial FOV 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 (to avoid the exact center of the scanner where the sampling density of lines of response may be very high), (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 (ramp filter with a cutoff at the Nyquist frequency); 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 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 
2 times 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 its 2 nearest neighboring points. The radial, tangential, and axial resolutions for each radius, averaged over both axial positions, are reported.
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. 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. Here, we define the scatter fraction to be the ratio of scattered events to total events, which are measured at a sufficiently low counting rate that random coincidences, dead-time effects, and 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 (2.3-mm inner diameter, 3-mL fillable volume) 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 (i.e., nearest to the patient bed), because the measured result will depend on the relative orientation of the line source and the bed.
Data are sorted into a sinogram, which is the 2D projection space representation (projection ray vs. angle) of a transverse plane. Oblique LORs are assigned to the slice where the LOR crosses the scanner axis using single-slice rebinning (5). Because the transaxial physical FOV of scanners varies, a fixed diameter of 24 cm (4 cm larger than the phantom diameter) is used for the calculation of scatter, as was done in the NU 2-1994 standard. The sinogram profile is used to calculate the number of scatter events within the FOV and the number of trues within a 2-cm radius of the source. The scatter within the peak is estimated by assuming a constant background under the peak, the level of which is determined by the average of the intensities near the edge of the peak (at ±2 cm). The sinogram profile is analyzed as a function of angle and the results are averaged. The scatter fraction for each slice and the average of the slice scatter fractions are reported.
Sensitivity
The sensitivity of a scanner represents its ability to detect annihilation radiation. In the NU 2-2001 standard, the absolute sensitivity of a scanner (expressed as the rate of detected coincidence events in counts per second [cps] for a given source strength, expressed in MBq) is measured. The measurement technique is based on the method of Bailey et al. (6). 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 distributed 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 underrepresent the efficiency of the scanner. For other measurements (e.g., scatter), a 70-cm-long phantom is used, as a way to standardize the phantom. A 70-cm-long uniform cylinder, however, would be impractical to fill and to handle. For those reasons, it was decided to move from the historical but arbitrary method of measuring sensitivity with a cylindric phantom to one less affected by the self-attenuation of a uniform cylindric source.
A 70-cm-long plastic tube is filled with a known amount of radioactivity, sufficiently low that count losses and randoms are negligible. The tubing is encased in metal sleeves of varying thickness and imaged, suspended in the center of the transverse FOV. Alternatively, the radioactivity can be placed in the smallest metal sleeve, as long as it is well sealed. Consideration was given to the sleeve material; it was agreed that the metal does not need to be specified, because the measurement is self-consistent. Single-slice rebinning is used to assign oblique LORs to the slice where the LOR crosses the scanner axis. The counting rate in the absence of attenuation Ro is determined by fitting the measured counting rates Rj as a function of sleeve thickness Xj:
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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 (NEC) rate (7) 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.
The total counting rate within a 24-cm transverse FOV is determined as the activity decays. The background, resulting from randoms and scatter, is estimated as was done for the intrinsic scatter fraction measurement. The true event rate Rtrues is then determined by subtracting the randoms + scatter background from the total rate Rtotal. The randoms rate Rrandoms is calculated as:
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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 encountered in clinical imaging, is reflected by the bias with which the scanner reports counts.
The accuracy of corrections for dead-time losses and randoms is determined from the counting rate performance data, described above. 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:
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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 (3,8), 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 measured 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, or 5.3 kBq/mL). The hot spheres are sequentially filled with activity concentrations of 8 and 4 times that of the background (2 sequential acquisitions). 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 background with nonradioactive water. After 1 half-life has elapsed, another sample of background activity (measured at the same time as the initial aliquot) 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 (as defined above) 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.
The acquisition protocol being simulated is to scan an axial distance of 100 cm in 60 min. The data acquisition time is determined by considering the axial distance z that the bed is translated between positions in a whole-body study (typically less than the axial FOV). The acquisition time Tacq is calculated by:
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The data are reconstructed with all available corrections applied, using the standard reconstruction algorithm and usual parameters for whole-body studies. To quantitate the image quality, ROIs with diameters equal to the physical inner diameters of the spheres are drawn on the spheres and throughout the background. Twelve background ROIs are drawn in the central slice, as well as in slices ±10 mm and ±20 mm away, for a total of 60 background ROIs of each size. The coefficient of variation of the means in these background ROI values for each sphere size is taken as a measure of the background variability. The hot sphere contrast recovery coefficient (CRChot) is calculated as:
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In addition, the accuracy of the corrections for attenuation and scatter is assessed by drawing an ROI in the region of the lung insert. The residual error (Clung) is calculated for each slice as:
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RESULTS AND DISCUSSION |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODSRESULTS AND DISCUSSIONCONCLUSIONREFERENCES |
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The measured spatial resolution will depend on the acquisition and reconstruction sampling. Figure 1 shows the transverse resolution on an ADAC UGM C-PET scanner (ADAC UGM, Philadelphia, PA) (9) for various sampling schemes. The radial sampling ranged between 0.5 and 4.5 mm; this was achieved by varying the transverse FOV between 128 and 576 mm and the number of radial samples from 128 to 256. The transverse angular sampling was 0.94° or 1.88°. Image pixel sizes ranged between 0.5 and 4.0 mm. The resolution achieved under clinical protocols can be noticeably worse than the best possible value, obtained with 0.5-mm radial sampling and 0.5-mm pixels. For example, in clinical brain imaging on this scanner, 2-mm radial sampling and 2-mm image pixel size are used, degrading the resolution by >1 mm from the best achievable value.
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Another difference between the previous and updated standards is in the specification of the width of the profile drawn for the purposes of generating the 1D response functions. The profile width can affect the measured spatial resolution if the response is not a gaussian function. Transverse resolution values (FWHM and FWTM) for the ADAC UGM C-PET scanner are compared in Table 1 for profile widths of 2 x FWHM, as specified by the standard, and 1 mm (2 pixels). The data had 0.5-mm radial sampling, 0.5-mm pixels, and a 4-mm slice separation. The results are for the source at x = y = 0 cm; because this scanner has continuous sampling, as opposed to a discrete ring of crystals, the FWHM at y = 0 cm is the same as that at y = 1 cm. A narrow profile (1–2 pixels), in general, leads to a lower FWHM, whereas a wide profile, as specified by NU 2-2001, results in a higher FWHM because of a potentially more asymmetric PSF. However, the narrower profile can also lead to a more variable result, whereas the wider profile is less sensitive to the relative position of the point source with respect to a voxel. In selecting a width of 2 x FWHM, the task force determined that reproducibility of the result was of highest importance.
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The reader is cautioned that the sampling in the axial direction of ring-based PET scanners may not be sufficient to produce a meaningful measure of axial resolution with the new standard. The value for the FWHM measured at a single axial location with such a system can vary dramatically, depending on the location of the source with respect to the imaging slice. This effect is demonstrated for the General Electric (GE) Advance scanner (General Electric Medical Systems, Milwaukee, WI) operated in 3D mode (slice separation, 4.25 mm) in Table 2, where the results of 2 repeated measurements are shown (similar results were obtained for 2D mode). In both measurements, the source was centered using the laser alignment device. The axial resolution measured, averaged over the 2 axial positions A (Z = 0) and B (Z = 1/4 FOV) specified by the NU-2 2001 standard, was 7.3 and 8.4 mm for the 2 measurements. However, if additional measurements are made with the source moved one half of a slice separation (2 mm for this scanner) from the locations specified by the NU 2-2001 standard (positions C and D), and the data for all 4 source positions are averaged, the resulting value for the axial resolution is much more stable (6.8 vs. 7.2 mm). Because it is difficult in practice to achieve the necessary positioning accuracy to guarantee exact centering of the point source on a slice, it is suggested that measurements on ring-based scanners be taken with the source located as specified in the standard but also moved one half of a slice separation. The results can then be averaged over the 4 axial locations. This issue was avoided in the NU 2-1994 standard by measuring the axial profile width instead of the axial resolution for scanners with axial sampling inadequate to perform the axial resolution measurement. The axial profile width measurement involved acquiring many scans at fine axial intervals. In addition, the scan data were not reconstructed, so effects of reconstruction processing on axial spatial resolution could not be assessed. Because the measurement was time-consuming and did not include processing effects, it was decided to measure the axial resolution on all scanners, regardless of sampling.
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An important consideration in interpreting the results of the intrinsic scatter fraction measurement is that the scatter fraction, as measured with either the 19-cm-long or 70-cm-long cylinder, is not representative of a realistic head or body size. It is useful for standard evaluations among scanners of different configurations. However, in the effort to use 1 phantom, the diameter of both phantoms (20 cm) is a compromise between a brain and a body. This makes it difficult to predict the absolute level of scatter that will be encountered in patient imaging situations. However, the scatter fraction measured with the 20-cm-diameter phantom on a particular PET system will indicate the relative level of scatter measured in a patient study.
Sensitivity
The total sensitivity values for several scanners are shown in Table 4 for radial positions of 0 and 10 cm. Results are also given using the uniform cylinder, as specified in the NU 2-1994 standard. Because the source configurations and definitions of sensitivity differ significantly between the 2 standards, it is difficult to compare the results directly. However, the relative ordering of sensitivity values among scanners should be the same for the 2 measurements, although there will not be a single scaling factor between the results for the 2 standards. As an example, the sensitivity values following the NU 2-1994 standard for the ADAC MCD system, with a 37-cm axial FOV, are significantly lower than would be expected from the NU 2-2001 values because the axial FOV is much longer than the older phantom (length, 19 cm). Also, there is a radial dependence of the NU 2-2001 sensitivity values for the coincidence gamma camera that is almost absent in the 2 circular systems.
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Counting Rate Performance
Table 5 shows typical peak counting rates for a GE Advance scanner, operated in both 2D and 3D modes, for the NU 2-1994 and NU 2-2001 standards. There is a dramatic decrease in the peak true and NEC rates for 3D acquisitions following the new standard compared with those for the old standard. This is largely due to the increase in randoms and dead time (decreased trues) as well as the higher scatter fraction from the OFA associated with the longer phantom. There was inadequate activity in the line source to achieve either the peak true or peak NEC rates in 2D mode on this scanner.
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A concern with using the line source to measure the counting rate performance was whether it yields counting rates (and effective activity concentrations) that are clinically relevant. Figure 2 shows a plot of the prompt rate as a function of singles rate for phantom and patient data. On the top are shown the counting rates for the ADAC UGM C-PET scanner for the 70-cm-long line source in a 20-cm-diameter polyethylene cylinder (following the NU 2-2001 standard) and whole-body patient studies. On the bottom are shown the counting rates for the University of Pennsylvania GSO brain scanner (G-PET) for the 20-cm-diameter x 19-cm long water-filled cylinder (following the NU 2-1994 standard) and brain patient studies. There is good correlation between the measured counting rates for the 70-cm line source and whole-body studies. Such a correlation has not been observed for the 19-cm-long uniform cylinder (9). Points lying below the phantom line correspond to patients whose attenuation is greater than that of a 20-cm-diameter cylinder. There is also good agreement between the measured counting rates for the 20-cm cylinder and brain studies. These results indicate that the 70-cm line source in the 20-cm-diameter polyethylene cylinder is a reasonable approximation to the clinical whole-body activity distribution (albeit for thin patients), whereas the 20 x 19 cm long cylinder is a good representation of the clinical brain activity distribution.
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10–15 mCi). However, for 2D systems with septa, the amount of activity required to reach the peak trues or peak NEC counting rates can be significantly higher (on the order of 5.5–7.5 GBq [150–200 mCi]); as has been seen, it can be difficult to achieve this activity level in the small volume of the line source. For 2D systems with septa, it may not be possible to measure the peak true or peak NEC counting rates with the line source method. It should be noted that these activity levels are generally not clinically relevant in whole-body FDG studies (where the activity concentrations are on the order of 4 kBq/mL) because of limitations on patient radiation exposure. However, the line source phantom may not be helpful at predicting the performance for bolus studies in 2D mode. An alternate approach would be to perform the measurement with the 20-cm-diameter cylindric phantom, following the NU 2-1994 standard. Because the effects of OFA are small for scanners operating in 2D mode with septa, the counting rate performance should be comparable to that achieved with high activity levels in the line source, following the NU 2-2001 standard. In addition, counting rates in brain studies correlate well with those seen with the 20-cm phantom (9), as seen in Figure 2. The NEC rate gives a measure of the signal-to-noise ratio (SNR) in an image, which could be used to assess the image quality at different activity levels. However, the NEC is only a global measure that is not sensitive to local variations of the SNR for particular source distributions. Because local variations (e.g., from event mispositioning) can occur at higher counting rates, it is important to look at the reconstructed images as well.
This measurement does not provide data for a realistic distribution, such as the brain or heart. The IEC group has suggested collecting counting rate performance data for 3 different configurations to simulate the head, heart, and thorax (3). The difficulty with this procedure is that 3 times as much data must be taken and processed. In addition, it has been our experience (9) that the IEC counting rate measurements for 3D systems do not, in fact, correlate well with cardiac and whole-body FDG studies because of the absence of activity outside the scanner. On the other hand, the counting rates for the 70-cm-long sources are well correlated with the counting rates seen in whole-body patient studies (9), as seen in Figure 2.
Following the NU 2-2001 standard, random events are estimated, rather than being measured separately (or calculated from the single event rates), as was assumed in the NU 2-1994 standard. The measurement of randoms is, therefore, independent of the scanner’s technique for determining the random event rate. The rationale for this change was that not all PET scanners currently have the ability to measure random coincidences accurately. However, there are several limitations to the method specified to estimate the scattered and random event rates. These background events are estimated by assuming a constant background under the peak in the sinogram, the level of which is determined by the intensities at ±2 cm from the peak. However, this background may include other misplaced events as part of the randoms measurement, especially at high counting rates. In addition, shifts in the photopeak with counting rate or any difference in dead-time effects for true, random, and scattered events may influence the randoms estimate, because the scatter fraction is assumed to be independent of counting rate. Finally, the assumption of a constant background under the peak in the sinogram ignores the spatial variation of scatter, so small-angle scatter is effectively included in the determination of true events.
It is important in this measurement that the activity in the line source be accurately determined. There are 2 possible methods by which this can be accomplished. The line source can be filled, coiled, and placed in a dose calibrator. Alternatively, the line source can be filled, and any residual activity remaining measured, with the line source activity being calculated from the difference of the 2 measured activities. The first has the advantage that the line source activity is measured directly; however, manipulating the line source in this manner can make it prone to leaks, and the response of the dose calibrator is sensitive to the geometry of the coil inside the well. The second method is straightforward, but care should be taken when filling the line source to ensure that no activity is spilled.
To compare the counting rate performance of various PET systems, it is necessary to specify the level of activity to which the scanner is exposed. The total activity in the line source is not a meaningful number if the FOV of the scanner is significantly less than the length of the source (70 cm), because not all of that activity is detected by the scanner. The activity inside the FOV could be an appropriate parameter but, for comparisons with other scanners, an activity concentration seemed more relevant. For that reason, the effective activity concentration is used as the common denominator in this measurement.
Accuracy of Corrections for Count Losses and Randoms
The results of measurements of the accuracy of corrections for count losses and randoms have been reported for 2 different scanners (9,11). For both systems, a rather large error (15%–40%) was reported at or below activity levels where the peak NEC rate is achieved. The use of a line source for this measurement, rather than a distributed source, may prove to be more challenging, from the point of view of accurate corrections for random coincidences and dead time. However, it should be noted that the activity levels where the NEC rate peaks may also be higher than those seen clinically.
The large ROI centered on the phantom was chosen rather than a smaller one, centered on the line source itself, because random (and scattered) coincidences are globally distributed throughout the FOV, not just in the region of radioactivity. Without proper correction for randoms throughout the image, image contrast will change as a function of activity level. Using a larger ROI permits an assessment of the global accuracy of the randoms subtraction.
Image Quality Measurement
Figure 3 shows a representative image for a 6-min emission scan from the study with an 8:1 ratio of sphere-to-background activities taken on the ADAC UGM C-PET scanner. Also shown is the placement of ROIs for the quantitative analysis.
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Because of variations in the uptake of radiopharmaceuticals and in patient sizes and shapes, it is difficult to simulate clinical imaging conditions using a single phantom. For that reason, the results of the image quality measurement should be interpreted with caution. For example, as has been noted (11), the walls of the hot spheres add a cold rim around the hot region that can decrease the measured hot sphere contrast.
The image quality measurement is intended to emulate whole-body imaging and, as such, would be less relevant for a scanner that is used primarily or exclusively for brain imaging. An alternative test for such a scanner, particularly one with a patient aperture smaller than 35 cm, would be to use the same hot and cold spheres in a 20-cm-diameter cylinder.
General Comments
For all measurements described except spatial resolution and image quality, a 70-cm long line source is used as part of the activity distribution. Another option would have been to tailor the length of the phantom to the axial FOV of the scanner. However, it was decided that a standard length was preferable because it permits the use of a single phantom on all scanners. In addition, the longer phantom provides a more clinically reasonable distribution of activity outside the scanner. This OFA can impact many of the performance measurements, especially scatter and counting rate performance.
Although the revised standard attempts to reflect PET scanner performance under conditions of whole-body FDG scanning, it is very difficult to mimic a clinical scanning situation with phantoms. The new standard does ensure that different scanners can be compared with each other. The relative results between scanners can be expected to hold for clinical situations as well. However, for all measurements, the results should be interpreted with caution. For example, the spatial resolution results may not reflect the clinical resolution achievable in patient images. As another example, the intrinsic scatter fraction measured with a 20-cm-diameter object will be lower than the scatter fraction seen in a body. Third, the counting rate performance measurement better predicts the trues and NEC rates for thin patients. Finally, the results of the image quality performance test depend on the reconstruction algorithm, as has been demonstrated, as well as the various corrections applied to the data. It is, therefore, not a static property of the scanner but, rather, a characteristic that can change when improved algorithms are implemented by the manufacturer. The measurements reflect the performance of a scanner under specific imaging conditions that may or may not reflect the actual imaging situation of a patient study.
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CONCLUSION |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODSRESULTS AND DISCUSSIONCONCLUSIONREFERENCES |
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ACKNOWLEDGMENTS |
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FOOTNOTES |
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For correspondence or reprints contact: Margaret E. Daube-Witherspoon, PhD, 11316 Smoke Rise Ct., Fairfax Station, VA 22039.
E-mail: daubew{at}bellatlantic.net
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODSRESULTS AND DISCUSSIONCONCLUSIONREFERENCES |
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J. S. Kim, J. S. Lee, K. C. Im, S. J. Kim, S.-Y. Kim, D. S. Lee, and D. H. Moon Performance Measurement of the microPET Focus 120 Scanner J. Nucl. Med., September 1, 2007; 48(9): 1527 - 1535. [Abstract] [Full Text] [PDF]
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C. C. Watson, M. E. Casey, B. Bendriem, J. P. Carney, D. W. Townsend, S. Eberl, S. Meikle, and F. P. DiFilippo Optimizing Injected Dose in Clinical PET by Accurately Modeling the Counting-Rate Response Functions Specific to Individual Patient Scans J. Nucl. Med., November 1, 2005; 46(11): 1825 - 1834. [Abstract] [Full Text] [PDF]
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S. Surti and J. S. Karp Imaging Characteristics of a 3-Dimensional GSO Whole-Body PET Camera J. Nucl. Med., June 1, 2004; 45(6): 1040 - 1049. [Abstract] [Full Text] [PDF]
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Y. E. Erdi, S. A. Nehmeh, T. Mulnix, J. L. Humm, and C. C. Watson PET Performance Measurements for an LSO-Based Combined PET/CT Scanner Using the National Electrical Manufacturers Association NU 2-2001 Standard J. Nucl. Med., May 1, 2004; 45(5): 813 - 821. [Abstract] [Full Text] [PDF]
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C. C. Watson, M. E. Casey, L. Eriksson, T. Mulnix, D. Adams, and B. Bendriem NEMA NU 2 Performance Tests for Scanners with Intrinsic Radioactivity J. Nucl. Med., May 1, 2004; 45(5): 822 - 826. [Abstract] [Full Text] [PDF]
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J. S. Karp, S. Su |
) taken on ADAC UGM C-PET scanner. (Bottom) Prompt rate vs. singles rate for 20-cm-diameter x 19-cm-long water-filled cylinder (solid line) and for brain patient studies (
