European Nuclear Medicine Guide
Positron emission tomography (PET) has been successfully used in cardiac, neurological and oncological research applications since the 1970s and numerous milestone publications have shown the potential of this technique for the characterization of functional tissue properties covering metabolism and perfusion. Since then, PET has been developed into a clinical imaging tool, but, due to high costs and lack of reimbursement, in many countries, PET stayed and still stays in the research domain. The driving factor for its success in the last decades was in 2001 when Townsend and Cherry implemented the first hybrid PET/CT system, which combined high-resolution morphological CT images with high physiological specificity of PET, albeit at a lower spatial resolution [1]. From a logistical point of view, this hybrid solution had another advantage which has to be appreciated: the time-consuming acquisition of a transmission dataset for attenuation correction using external sources, taking up to 15 minutes per bed position, was replaced with a CT scan lasting a few seconds. This enabled a higher patient throughput in whole-body, oncological imaging and paved the way for the success of PET/CT in today’s medical imaging world. Although this success is almost entirely based on tumor imaging, with the clinical availability of PET/CT systems neurological and cardiac applications are still widely used (Fig. 1).
Figure 1: Three PET/CT examples from oncological imaging, neuro-oncology, and cardiovascular protocols.
PET imaging uses the radioactive decay of positron-emitting radionuclides (such as 18F, 68Ga, 82Rb, 11C, 13N, etc.), which are, after their generation in cyclotrons or elution generators, chemically “inserted” in molecules with high biological relevance (e.g. for assessment of metabolism, receptors, and perfusion). These radiopharmaceuticals (otherwise called tracers) are intravenously injected into a patient either prior to the scan or after positioning in a scanner (Fig. 1). The emitted positrons will annihilate with electrons within a very short time under the emission of two 511 keV photons (this process is discussed in more detail earlier in this guide. In contrast to SPECT, where collimators are used to generate the geometrical information, PET/CT tomographs utilize the near-simultaneous (coincidence) detection of these photons with rings of detectors, the so-called electronic collimation, with a much higher efficiency as compared to SPECT (Fig. 2). The high-energy photons are detected with dedicated crystal materials which convert them into lower-energy photons, which are then, in turn, detected and amplified with mostly digital detectors. These electrical signals are subsequently processed to estimate the so-called line of response (LOR) along which the decay must have occurred. This is possible since the two annihilation photons travel in an anti-parallel manner (i.e., they are emitted at an angle of approximately 180o), implying that the decay must have taken place somewhere along the LOR (Fig. 3). Thus, PET relies on a coincidence detection by using a very short-acceptance timing window of typically a few nanoseconds or less for the detection of both events. Fundamentally, three types of events can occur (Fig. 4). The first kind is a true coincidence, if indeed the two annihilation photons arrive unscattered at the two detectors along the LOR. The second type of event, a scattered coincidence, occurs when one or both photons from a single positron decay undergo a scatter event in the body, but arrive within the time window; obviously, this results in a wrong spatial association. Finally, a random coincidence can occur, where rays from two unrelated decay events are registered within the time window. Whereas the first possibility results in a “correct” measurement, the two other cases will yield image degradation. Originally, PET systems operated in a 2-dimensional (2D) mode with interplane septa, which reduce scattered photons in coincidence measurements. Basically, in this 2D mode, coincidence measurements are performed only in one plane of the PET camera. Removing the septa and accepting coincidences between scanner planes, however, increases sensitivity and consequentially shortens scan time. This so-called 3D mode has shown many advantages especially in oncological imaging protocols, including increased patient throughput [2, 3]. PET systems using 2D mode belong to legacy and are long gone, the technique is mentioned here to start out where it all began.
Today, the most commonly used detector materials are lutetium oxyorthosilicate (LSO) and gadolinium oxyorthosilicate (GSO), both attractive due to their physical properties. LSO and GSO are increasingly used instead of bismuth germanate (BGO) in PET/CT systems [4]. Their relatively fast light decay time and high light yield [5] enable the use of short coincidence time windows. Consequently, this improves the count rate capabilities and reduces randoms. With the aim to improve spatial resolution, recent PET/CT systems are equipped with smaller crystals, which is of particular importance as it potentially allows the detection of smaller structures. As the enhancement of spatial resolution from originally 7.0 mm to 3.2 mm leads to a significant increase in count recovery, the use of high-resolution PET ameliorates the assessment of regional tracer distribution in the targets and allows for more accurate quantification of physiological parameters such as metabolism, receptor density, or blood flow.
From the CT components perspective, the available options range from 16-slice to 128-slice systems, depending on the desired application. This corresponds to the variety of protocols found in (typically) oncological clinical scenarios. This spectrum ranges from very high-throughput centers with low dose CT protocols to sites where a full contrast media enhanced diagnostic CT workup is combined with PET. The dominating metabolic tracer used worldwide is clearly 18F-FDG but the use of approved radiopharmaceuticals for neuroendocrine and prostate cancer is rapidly increasing.
Figure 2. Basic principles of PET imaging: a radioactive tracer is injected into a patient and (B) emits two 511keV photons after the positron annihilates with an electron. Using a ring design, photons are detected and measured with photomultiplier tubes (PMTs) surrounding the patient.
Figure 3. Comparison of PMT- (A) and SiPM- (B) based detector modules (figure adapted from [add reference https://doi.org/10.1201/9780429489556-22]
Figure 4. Millions of signals per second coming from all the PMTs are processed in coincidence mode: only those two events that are detected within a few nanoseconds (or less for time-of-flight) are considered to stem from the same annihilation event [24].
Figure 4. Three possibilities exist for such a measurement: true, scattered and random coincidences. The true events are identified using sophisticated algorithms
Attenuation correction in PET is the prerequisite for any quantification of the radiotracer uptake signal (Fig. 5, 6). Such an absolute quantification in units of Bq/ml is the key to superior diagnostic performance without attenuation artefacts, its normalization to the injected dose and the patient weight (SUV) enabling comparisons between serial examinations and performing any pharmacokinetic modeling. A large fraction of the 511 keV annihilation photons from the positron decay are actually scattered by the patient’s body. Consequently, they are discriminated against due to a lower energy or do not reach the PET detectors at all. To account for these effects and thus compute activity-wise correct PET images, it is necessary to determine an attenuation map with the appropriate attenuation coefficients for 511 keV photons at each voxel. In hybrid PET/CT systems, this is achieved by using the information about the tissue electron density, provided by the CT, and adjusting it with regard to the difference in photon energy. However, although the CT scan is very fast and PET scan times are being constantly reduced, misalignment can still occur: misregistration between emission and “transmission” might lead to uptake errors requiring strict quality control.
In addition to the effects of misalignment, metal implants, or other interventional devices may affect the quantification of the PET tracer uptake. These PET artifacts are primarily due to the reconstruction artifacts of the CT and migrate through the overestimating of attenuating tissue into the PET images. The relevance of these artifacts differs depending on their position; thus, hybrid reading, including the review of non-attenuation-corrected data, is advisable.
Figure 5. The most relevant correction in PET is attenuation correction (AC). These examples of a cylindrical phantom shows the significant effect of attenuation towards the inner portion both in the images and the line profiles. Note, that the signal fluctuations in the non AC data is enhanced – as this is what is actually measured by the scanner.
Figure 6: As was shown in Fig. 5 and a phantom, the signal loss towards the objects center in non AC data shows the liver uptake variable with the location. This is recovered when attenuation correction data from the CT is included in the image reconstruction process.
Since the introduction of PET/CT at the turn of the millennium, a series of technical improvements have happened. From a more logistical aspect, many (but not all) developments of the stand-alone CT systems were made available for the “hybrid cousin”. Today, with the exception of rather dedicated, high-resolution cardiac CT systems, a variety of CT components in a PET/CT cover the area from rather cost-effective systems to the high end.
On the PET side, more intrinsic developments have occurred in the detector, acquisition, and image reconstruction domains.
Utilizing the time-of-flight information is actually not a new concept, but was described already in the early days of PET [6]. As outlined before, PET scanners utilize a short temporal coincidence window to decide whether the detected photons originate from the same annihilation event. With improved detector technology, the temporal resolution of PET systems enabled the measurement of the difference in the arrival time of these photons, which, in turn, allowed the estimation of where the event happened along the LOR [7]. Since 2006, all major vendors have implemented this technology in their systems, which show, especially in obese patients, improved image quality and thus better lesion detectability.
To the same extent the PET detector technology improved, computer processing power did too. This resulted from a complete migration from analytical reconstruction techniques (“filtered back projection”) to non-analytical approaches, where the particular image is estimated such that its projection data optimally fits the acquired projection data. As this also allowed the integration of CT data for attenuation and scatter correction, it paved the way for dramatically improved image quality [8]. This led to a long series of incremental improvements in the PET image reconstruction, where many factors describing the properties of the imaging apparatus are integrated into the reconstruction algorithms [9]. This approach is known as point-spread-function modeling or resolution recovery. It is worth noting that all these complex computations impact quantification and require harmonization, especially with respect to serial examination and multi-center studies [10-12].
Even though cardiac PET gating has been clinically implemented since decades, the use of respiratory gating in oncological imaging is still less widespread. One contributing factor to this effect is the acquisition duration relative to the object size: in cardiac PET, 10-20 minutes are used to cover the heart in a single bed position, whereas in whole-body oncological PET imaging, scan times of 1-3 minutes in the thorax simply result in less data.
However, respiratory motion during the acquisition interval discernibly blurs the signal from the affected structure and thus reduces the spatial resolution that could theoretically be achieved. Technically, any respiratory and/or cardiac gated PET uses an implementation where, in parallel to the measurement of all coincidence events, the gating signals are recorded and stored together in a so-called list-mode stream [12]. For cardiac gating, an ECG is used, and for respiratory gating, a pneumatic device is typically integrated into an elastic belt, which is then fastened around the lower chest of the patient. This allows for the association of a given annihilation event with the motion state and enables sophisticated approaches in motion correction. From a technical perspective, it is worth mentioning that the frequency distribution of human respiration is significantly different to that of cardiac gating as the respiratory frequency is much more irregular. Thus, its distribution may vary substantially over the length of a PET acquisition as patients might go from anxious into a more relaxed state or even fall asleep. A related improvement is the continuous bed motion. Historically, when PET examinations required only coverage of the axial extent of the patient for cardiac and neuroimaging applications, an extended scanning range was realized with several, overlapping bed positions. Technically, the approach of a continuously moving bed (similar to CT and MRI) enables - together with adaptive table speeds - a more homogeneous sensitivity along the scanner axis, while also increasing patient comfort as the more or less abrupt change from one bed position to the next is avoided [13].
The final major technical improvement was the replacement of the conventional photomultiplier tubes (PMTs) in PET systems which recently entered the clinical arena. Initially developed for PET/MR systems, which do not allow for conventional PMTs, the new (“digital PET”) devices, based on avalanche photodiodes (APD) [14] and silicon photomultipliers (SiPM) [15], show significantly increased sensitivity and also allow for a more compact design. The improved image quality can be utilized in several ways: higher detection rates, reduced scan times, dynamic imaging or lower radiation doses. How this increased flexibility will be used in which settings is still an open question, but it provides another step in our journey optimizing personalized medicine [16]. However, with respect to reliable quantification, the introduction of new hardware and algorithms has increased the need for harmonization of quantification and remains an important element [17].
In addition, these approaches offer the (albeit rather expensive) means to implement large-field-of-view (LaFOV) scanners with simultaneous scan lengths between one and two meters [18 - 20] enabling imaging either at low to very low amounts of activity applied or shorter to very short scan durations as well as dynamic imaging further improving our understanding of complex physiological processes simultaneously throughout the whole body [21]. For Fluor-18-labelded tracers, this results in either very low radiation exposures on the order of the annual natural exposure (1-2mSv) per procedure and, due to scan times of only several minutes well below 5 for the whole body, much higher compliance of patients with the imaging procedure, even at levels severe morbidity. However, currently one could state that the situation is similar to the early PET/MRI days where the number of published review articles was greater than the number of actual research papers.
As in all other areas of our field, the use of artificial intelligence is potentially offering improvements in signal processing, image reconstruction and data processing. Examples range from its use in image reconstruction [22], improved analysis [23] and tumor load assessments [24] - an elsewise very tedious manual process - just to name a few of the applications that will be available, although the regulatory burden is high compared to conventional approaches.
Without any doubt, molecular imaging using PET/CT is an invaluable tool in clinical routine. As of today, more than 1000 PET/CT systems are installed in Europe alone. Worldwide installed are more than 3000 PET/CT systems, amongst those with a long axial field of view (LAFOV-PET/CT) currently count about 50, further increasing due to the aforementioned advantages with respect to patient comfort, -throughput and availability.
Although still a technically ambitious and logistically challenging methodology, it provides an unprecedented level of performance in terms of sensitivity, specificity, and scan time with continuous improvements shaping our path from physics to physiology.
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An extensive collection of PET/CT artefacts can be found in PET/CT Atlas of quality control and image artefacts. IAEA Human Health Series no. 27, IAEA, Vienna 2014
https://www.iaea.org/publications/10424/pet/ct-atlas-on-quality-control-and-image-artefacts (accessed 2020 05 07)