I. INTRODUCTION
3. PET imaging of CXCR4 expression
Positron emission tomography (PET) utilizes the decay of positron-emitting radionuclides in which a proton in the nucleus is transformed into a neutron, a positon (β+) and a neutrino (ѵ).
After ejection from the nucleus, the positron loses its kinetic energy through interaction (collision, ionization or electronic interaction) with atoms of the surrounding matter and comes to rest, usually within a few millimeters of the site of its origin in body tissues (depending on the energy of the ejected positron and the type of surrounding matter). The positron and an ordinary electron temporary form a “pseudo-atom” called positronium, which has a mean lifetime of 125 picoseconds and is converted according to the mass-energy equivalence (E = (m+
+ me) • c2) into two 511 keV annihilationphotons (ɣ-photons) that are emitted in mutually opposite directions, see Figure 6 105. A selection of positron emitting radioisotopes used in nuclear medicine, their half-lives and maximal positron-energies is depicted in Table 2.
Near-simultaneous detection of the two annihilation photons allows PET to localize the origin of positron-emitters along a line between the detectors (Fig. 6). This mechanism is called
17 annihilation coincidence detection (ACD) and is usually measured within a timing window of 6 - 12 nanoseconds and an energy window of 300 – 650 keV. By incorporation of multiple opposing detectors in a complete ring around the patient, data for multiple projection angles can be acquired simultaneously. However, the quality of an image produced by PET is degraded by several physical factors like detector sensitivity differences, random coincidences or scattering.
Table 2. Physical properties of selected PET isotopes (positron emitters) 106.
radioisotope half-life (t1/2)
maximum β+ -energy (abundance)
11C 20.4 min 1.0 MeV (99.8%)
13N 10.0 min 1.2 MeV (100%)
15O 2.0 min 1.7 MeV (99.8%)
18F 109.7 min 0.6 MeV (96.9%)
64Cu 12.7 h 0.7 MeV (19.3%)
68Ga 67.6 min 1.9 MeV (90%)
89Zr 78.4 h 0.9 MeV (22.7%)
124I 4.2 d 2.1 MeV (25%)
Therefore flat, ring-shaped lead or tungsten septa are used, not only to reduce the number of scattered events collected, but to minimize other effects of radiation originating outside the field of view 107.
To further optimize the PET signal, corrections for attenuation, dead-time and pile-up events need to be applied to the projections prior to reconstruction. Finally, algorithms for reconstructing PET images such as filtered back projection (FBP) and ordered subsets expectation maximization (OSEM) are used to process the raw data (sinograms) 108. The images that result from PET provide quantitative information about the voxel intensity and the amount of radioactivity in a voxel. Calibration factors must be determined to translate the corrected counts to radioactivity values (kBq/cm3) 105. The signal in a region of interest (ROI) can then be
18 described as % injected dose per mL (% ID/mL), activity per volume (Bq/cm3), or as standar-dized uptake value (SUV), which is defined as the activity concentration in thevolume of interest [kBq/mL] times the body weight [kg] divided by the injected activity [kBq] 109.
Figure 6. Schematic representation of a PET scanner. The radioisotope decays by β+-emission.
Subsequent annihilation of the formed positronium results in two 511 keV ɣ-photons, which are counted by two opposite detector units electronically connected via a coincidence circuit.
If signals originate from very small structures, they will have their radioactivity concentrations either overestimated or underestimated, since the activity signal (same total counts) seems to be distributed over a larger volume due to image blurring and the way of image sampling during the PET signal analysis (partial-volume effect) 105, 110. In addition, the finite positron range (energy of ejected positron) and photon noncollinearity (annihilation photon are not exactly 180°
apart) also contribute to the spatial resolution of the resulting PET images. The resolution of a preclinical scanner can reach up to 2.5 mm and 4 – 6 mm on whole-body PET systems, respectively 105.
The in vivo behavior of a tracer is varying with time and depends on a number of components.
Tracer delivery, extraction from the vasculature, diffusion or transport into cells, metabolism and excretion from the body (also referred to as the ADME principle in pharmacokinetics with
19 absorption, distribution, metabolism and excretion). Dynamic PET imaging allows direct measurement of the radioactivity concentration over different time frames and can therefore be employed to describe the kinetic of a radiopharmaceutical in the body 105, 111. By applying ROIs on different compartments (heart for blood pool and tumor for specific binding site for example), the dynamic change of tracer concentration in these compartments can be observed and support the understanding of tracer distribution in vivo.
CXCR4 ligands for PET: As a small-molecule CXCR4 antagonist used for PET imaging of CXCR4, [64Cu]AMD3100 appeared to be fast and efficiently radiolabeled and showed rapid clear-ance from the blood and accumulation in CXCR4 specific tissue. However, due to high accumu-lation in the liver (>40% ID/g), the clinical applicability of this tracer is challenging 112, 113. An optimized AMD3100 derivative, AMD3465 revealed promising target properties with very high accumulation in CXCR4+ tumors (> 100% ID/g), but also significant accumulation in the liver (40% ID/g), whereas the 11C-labeled analogue unfortunately exhibited low tumor-to-background ratios 114, 115. Other small-molecule CXCR4 antagonist [18F]MSX-122F and [18F]M508F (see Figure 7) displayed specific binding to CXCR4 in vitro, but were not further evaluated 116, 117. Radiolabeled analogues of CXCL12 (125I and 99mTc) were only used for in vivo biodistribution, but the specificity of the signal is doubted, due to rapid enzymatic degradation 118, 119. In addition 125I-labeled antibodies for CXCR4 were used for in vitro and in vivo biodistribution studies, but were not able to clearly distinguish CXCR4+ from CXCR4- tissue 120. Very recently an 89Zr-labeled human CXCR4-mAb (89Zr-CXCR4-mAb) was evaluated for detection of CXCR4 expression. In vitro and in vivo evaluation of 89Zr-CXCR4-mAb showed enhanced uptake in CXCR4+ xenografts.
20 Figure 7. Structures of selected PET imaging agents for CXCR4 targeting.
It also demonstrated the ability to detect lymph node metastases in an experimental model of metastatic triple negative breast cancer. However, due to slow antibody clearance kinetics, late imaging time points (optimum 7 days p.i.), and thus somewhat complicated imaging protocols, the clinical applicability of 89Zr-CXCR4-mAb is also challenging 121.
Radiolabeled derivatives of T140 (see Figure 7) were also employed for PET imaging. 4-18F-T140 showed low tumor-to-background ratios mostly due to enhanced binding to red blood cells and accessory elevated uptake in liver tissue 96. Even though the exchange of the fluorobenzyl group with DOTA or NOTA reduced the unspecific binding to red blood cells in 64Cu-DOTA-NFB 122,
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64Cu-NOTA-NFB 122 and Al[18F]NOTA-T140 95, a significant accumulation in liver tissue remained, which resulted in low tumor-to-background ratios. First clinical application was reported for 68Ga-NOTA-NFB in healthy volunteers and glioma patients. Good tumor-to-background ratios and a low tumor-to-background uptake were reported. However, 68Ga-NOTA-NFB primarily accumulates in the spleen and the liver, which resulted in a slightly higher effective radiation dose compared to [68Ga]Pentixafor 123. The 68Ga-labeled T140 derivative 68Ga–CCIC16 demonstrated favorable pharmacokinetic properties along with CXCR4 specific accumulation (tumor-to-muscle ratio: 9.5) 94.
Radioiodinated FC131 (R2) was the first cyclic pentapeptide based imaging agent. Unfortu-nately, 124I-FC131 is very lipophilic, which is thought to be responsible for high uptake in the liver and intestines, as lipophilic compounds are often excreted via the hepatobiliary route (partition coefficient logP = −0.35 ± 0.02, as determined in octanol/PBS) 124. Intensive research and a library of peptide ligands resulted in [68Ga]Pentixafor ([68Ga]23, logP = −2.90 ± 0.08), that exploits a 4-aminomethyl-benzoic acid linked hydrophilic DOTA chelator for labeling. Due to its highly specific binding to human CXCR4 and favorable pharmacokinetics, [86Ga]Pentixafor is currently the only radiopharmaceutical suitable for CXCR4 imaging in patients and is assessed in a broad range of clinical proof-of-concept studies for a variety of diseases (cancer, cardio-vascular diseases, stroke or inflammation) 90, 91, 125-133.