**4. Discussion**

Interpretation of SUV metrics is a valuable tool in the assessment of PET/CT scans, as clinically relevant parameters such as d'Amico risk classification, PSA plasma levels and Gleason score correlate significantly with SUV [37–39]. However, SUV is also affected by aspects inherent to the imaging method such as uptake time [40], reconstruction algorithm used and the use of PSF modelling [41,42], bed motion [43], use of breathing instructions [44,45], scan time [46] and scanner properties [47]. Therefore, caution is warranted when interpreting SUV for clinical evaluation of 68Ga-PSMA PET/CT scans. Differences in pharmacokinetics and pharmacodynamics should be considered when comparing uptake values obtained from scans with different tracers.

Improved lesion conspicuity and increased SUVmax for Q.Clear reconstructions with low β are described in the literature [26]. Lowering the β corresponds to less noise suppression and therefore higher SUVmax values. For SUV measurements, low β values are found to be more accurate when considering the average uptake in a lesion.

This effect is noticed in phantom scans for measurements of the RCmax for both T/B ratios, all simulated acquisition times and all spheres considered in this study. As the RCavg is dependent on the maximum voxel value, this effect is also present in the average recovery curves but less pronounced due to averaging over a larger number of voxels. The RCs exhibited by PSF and OSEM reconstructions are affected by the 6.4 mm Gaussian post-filter, which was chosen based on clinical reconstruction settings in our institute. Lowering or eliminating post-filtering, RCs will increase. On the other hand, even with the post-filter applied, noise levels based on the background variability measurements are higher for PSF and OSEM reconstructions than for any of the BPL reconstructions considered.

The higher recovery coefficients measured for shortened acquisition times are consistent with the increase in SNR. The maximum voxel uptake value is likely to increase when

the number of counts is decreased, as the signal-to-noise ratio is proportional to the square root of the number of counts (9):

$$\frac{\text{Signal}}{\text{Noise}} \sim \sqrt{\text{N}} \tag{9}$$

Therefore, both the average and the maximum apparent recovery coefficient increase when the number of counts taken into account in the reconstruction is decreased. This effect is less pronounced with increased β, due to the smaller noise tolerance and therefore smoother images from high β reconstructions. In general, caution is needed when comparing SUVs between two scans in which administered activity or scan times differ.

As the two phantoms used in this study were scanned simultaneously, acquisition of the bed position containing the spheres in the Micro Hollow Sphere phantom was started 10 min after acquisition of the bed position containing the spheres of the NEMA Image Quality phantom. Therefore, the activity concentrations in the Micro Hollow Sphere phantom were approximately 6% lower than those in the NEMA Image Quality phantom. The resulting decrease in the number of counts detected probably has a small effect on the maximum voxel value, and may contribute to the difference in recovery coefficients found in the NEMA Image Quality phantom and the Micro Hollow Sphere phantom.

Due to spill-out, RCs are affected by lesion size for smaller lesions. Looking at the sphere diameter at which the spheres' RCavg deviates significantly from that of the larger spheres in the same reconstruction, a dependence on the β is noted. For higher β, the decrease in RC starts at larger diameters. The volume of each of the three smallest spheres considered in this article (33.51 mm3, 65.45 mm<sup>3</sup> and 113.1 mm3) is smaller than five voxels using the minimal voxel size of the used PET/CT scanner (24.37 mm3). Coincidental high count rates in a single voxel, for example induced by a coincidental centering of a voxel amid a sphere, can induce a 3D isocontour at 50% of the maximum voxel value that consists of a single voxel. This will result in a positive RC bias, an overestimation of the recovery coefficient.

A large increase in average recovery coefficient observed for the 8 mm-diameter sphere for T/B ratio 10:1 and the 6 mm sphere for T/B ratio 20:1, most evident at low β, is worth mentioning. Detailed inspection of the reconstructions revealed that these spheres appeared to be coincidentally aligned with the reconstruction matrix. As the diameter of the spheres is smaller than three times the minimum voxel dimension, the exact position of the phantom defines the number of voxels over which the total number of counts from the sphere are distributed and therefore strongly influences the recovery coefficient. The effect can be enhanced by a coincidental high number of counts due to Poisson noise, which means the effect is more likely to be noticed for lower β, shorter acquisition times and lower activity concentrations. Taking RCpeak rather than RCavg as a measure, the voxel sampling effects are eliminated leading to more robust results. However, as the 1.2 cm-diameter spherical VOI used for obtaining the RCpeak is larger than the hot spheres in the Micro Hollow Sphere phantom, this method incorporates background voxels in the VOI, leading to a lower RC. Therefore, in small lesions, SUVpeak cannot be used to discriminate between larger volumes with low uptake and smaller lesions with high uptake.

The findings from this study are comparable to those described in 18F-FDG PET/CT studies. Improving contrast recovery for lower noise penalties in BPL reconstructions is well described by Teoh et al. [28,29] and similarities between the preferred β values for patient scans in this study and those recently described by Messerli et al. for 18F-FDG are also noted [48]; the observation that voxel sampling influences measured uptake values is in line with results for 18F-FDG PET/CT shown by Mansor et al. [49] and the observation that RCs decrease for increasing T/B ratio is described by Munk et al. [50]. These similarities are explained by the fact that, from a physics point of view, the main potentially relevant difference between use of 68Ga and 18F is the positron range.

For a PET system, the spatial resolution can be written as (10):

$$\mathbf{R\_{sys}} \approx \sqrt{\mathbf{R\_{det}^2} + \mathbf{R\_{range}^2} + \mathbf{R\_{180}^2}}\tag{10}$$

where Rsys is the spatial resolution of the system, Rdet is the contribution of the detectors, Rrange is the contribution of the root mean square (RMS) positron range in water and *R180* is the contribution of the noncollinearity of the annihilation photons [51]. Assuming a system resolution for 18F of approximately 5 mm FWHM [15] and evaluating in the RMS positron ranges of 0.23 mm for 18F and 1.2 mm for 68Ga [52,53], it is evident that the increased positron range only yields an incremental increase in spatial resolution.

To summarize, comparison of SUV measures between different lesions or the same lesion in two different scans is not straightforward even when administration, scanning and reconstruction protocols are equal.

This finding is in line with the conclusion by previous authors that quantitative measures for small lesions in PSF reconstructed PET images can lead to misinterpretation as they vary with lesion size and are less reproducible [50].

Assessment of the reproducibility of RCavg and detectability of lesions in terms of the COV, RCavg and BV for different β suggests a value of 600 as an optimum when quantification as well as detection is of importance. Higher values yield impaired detectability as small lesions blur into the background. Lower values will lead to more accurate uptake measures and better detectability for small lesions. However, the introduction of additional noise will probably yield an increase in false-positives and lower reproducibility which is of particular importance for test–retest studies and follow-up scans.
