¹⁸F-fluorodeoxyglucose (FDG) is the most commonly used and the only oncologic PET tracer approved by the Food and Drug Administration (FDA) for routine clinical monitoring of tumor glucose metabolism . More than 90 % of oncologic PET imaging is performed by FDG-PET due to the increased metabolism of glucose by most of the solid tumors including the lung, colorectal, esophageal, stomach , head and neck, cervical, ovarian, breast, melanoma, and most types of lymphomas. In addition to diagnosis, staging, restaging, and monitoring response to cancer treatment, FDG-PET can be useful for selection or delineation of radiotherapy target volumes. FDG-PET has been used as a dose-painting target for sub-volume boosting and thus guiding radiotherapy planning. The use of FDG-PET for radiotherapy planning purposes has shown increasing importance, as more and more radiation oncologists believe that target volume selection and delineation can be adequately performed using FDG-PET in certain cancer types such as non-small cell lung carcinoma (NSCLC) stage N2-N3 patients [53]. In head and neck squamous cell carcinoma (HNSCC) , methodological studies have shown that the use of pre-radiotherapy using FDG-PET led to a better estimate of the exact tumor volume, as defined by the pathologic specimens, compared with CT and MRI. Interestingly, when validated segmentation tools were used, the mean FDG-PET-based GTV was consistently smaller than the GTV defined from morphologic imaging in all investigated tumor locations and at all-time points during radiotherapy (Fig. 5.2a) [12, 56, 57]. ¹⁸F-FDG PET/CT is also increasingly used for monitoring tumor response after completion of therapy as demonstrated for malignant lymphoma lung, colon, and breast cancer [58]. An example from a lymphoma patient shown in Fig. 5.2b demonstrates that high metabolic activity (measured by high uptake of FDG) of the tumor before the therapy had reduced after the therapy (indicated by reduced FDG uptake), indicating the efficacy of the treatment [54]. Persistently increased FDG uptake after treatment is also associated in predicting a high risk for early disease recurrence and poor prognosis. In patients with Hodgkin’s disease and aggressive non-Hodgkin’s lymphoma , ¹⁸F-FDG PET showed very promising results for assessing tumor response early in the course of therapy. This study included 260 patients with Hodgkin’s lymphoma and utilized ¹⁸F-FDG PET before and after two cycles of chemotherapy to monitor tumor response. The 2-year progression-free survival for patients with positive PET results after two cycles of chemotherapy was 13 %, whereas it was 95 % for patients with a negative PET scan. In a univariate analysis, the treatment outcome was significantly associated with PET response after two cycles of chemotherapy and various well-known clinical prognostic factors such as stage and the international prognostic score (IPS) . In multivariate analyses, however, only positive PET results after two cycles of chemotherapy turned out to be significantly correlated with patient survival (P < 0.0001). These data indicated that tumor response in ¹⁸F-FDG PET after two cycles of chemotherapy is a stronger predictor of patient outcome than the IPS and other well-established clinical prognostic factors [59]. The results from this study concluded that ¹⁸F-FDG PET appears to be the single most important tool for risk-adapted treatment in Hodgkin’s lymphoma [59]. The ability of ¹⁸F-FDG PET to predict tumor response early in the course of therapy as in the case of Hodgkin’s lymphoma offers the opportunity to intensify treatment in patients who are unlikely to respond to first-line chemotherapy. Conversely, treatment could be shortened in patients who show a favorable response after two cycles of chemotherapy . This is of particular interest in Hodgkin’s lymphoma, since chemotherapy combined with radiotherapy can cure most of the patients but also puts them at increased risk for secondary malignancies and other serious long-term complications, such as infertility and cardiopulmonary toxicity . Focal ¹⁸F-FDG uptake after chemo- or radiotherapy has been shown to be a strong prognostic factor. In one of the largest prospective studies published so far, Schwarz et al. [55, 60] prospectively performed ¹⁸F-FDG PET in patients with cervical cancer treated by chemoradiotherapy . Post therapy ¹⁸F-FDG PET (2–4 months later) showed a complete metabolic response (Fig. 5.2c) in 70 % of the patients, a partial metabolic response in 16 %, and progressive disease in 13 % of patients. The 3-year progression-free survival rates of these patient subgroups were estimated to be 78 %, 33 %, and 0 %, respectively (P < 0.001) [55, 60]. Another application of ¹⁸F-FDG PET could be in deciding sub-volumes within tumor or lymph nodes that demonstrate high metabolic activity and therefore represent high-risk lesions to be selectively treated with an increased radiation dose. Several studies in rectal and lung cancer have shown that FDG-PET allows selective boosting of hypermetabolic areas. For example, in a study conducted in patients with small-cell lung cancer, FDG-PET-based radiation planning for mediastinal lymph nodes changed the radiotherapy field in 24 % of the patients [61]. In patients with head and neck cancer, the radiation boost dose was markedly elevated and directed at the tumors with the highest FDG-avidity, and the adverse treatment-related effects remained limited [62]. Despite its great clinical utility, FDG-PET has a few limitations. The technique has lower sensitivity for slow growing, metabolically less active tumors (such as prostate, thyroid, and neuroendocrine tumors), and high levels of uptake in some normal tissues, such as the brain, that can make quantification of tumor uptake specifically difficult. Accumulation of PET tracers in infiltrating inflammatory cells, which also show high glycolysis, might give false-positive results and limit the sensitivity of the technique for detecting tumor response to treatment [63].

Hyperpolarized ¹³C MRI is an emerging molecular imaging technique that can provide unprecedented gain in amplifying signal intensity. This technique can be used to monitor uptake and metabolism of ¹³C-labeled endogenous biomolecules such as glucose, pyruvate, fumarate, etc. [64, 65]. The degree and magnitude of the increase in sensitivity of biomolecules depend on the extent of polarization achieved, the T1 relaxation time of the agent, the delivery time, and the MR imaging methods employed. The polarization of ¹³C at a magnetic field strength of 3 T is calculated to be 2.5 ppm. Consequently, biological molecules enriched with ¹³C suffer from poor sensitivity because of a lower magnetic moment, lower polarization, and reduced concentrations compared to tissue water proton (in range of a few mM versus 80 M for water ¹H). As a result of these differences, ~ four orders of sensitivity difference for ¹³C-containing molecules compared to water protons need to be bridged to implement metabolic imaging using endogenous organic molecules. Increasing the polarization (fraction of molecules with nuclear spins emitting signal) needs to be increased significantly. Among the various methods used to enhance the polarization of nuclei, dynamic nuclear polarization (DNP) technique has so far been the most successful for in vivo applications [64–66]. In DNP, the higher polarization of a molecule with an unpaired electron spin is transferred to nuclei such as ¹H, ¹³C, ¹⁵N and ¹⁹F, etc. Firstly, for a molecule to be sufficiently polarized with long signal decay, it should possess a carbon site with long intrinsic T1 that can be enriched in ¹³C (e.g., carbonyl or carboxylic acid). Introducing deuterium in aliphatic carbons has been shown to be very amenable. Secondly, the molecule should be able to form a homogenous glassy solid formulation with the paramagnetic agent as when frozen it must ensure polarization transfer from the electrons to nuclei. Thirdly, upon dissolution the loss of polarization should be marginal and not significantly high. Fourthly, the chemical shifts of the injected molecule and products should be clearly distinct. Lastly, the molecule should not exhibit any associated side effects when injected as a bolus at the required doses (typically approx. 0.1 mmol/kg body weight). Pyruvate labeled with ¹³C in the C-1 position satisfies all the prerequisite conditions listed above for a polarized agent. Hyperpolarization of pyruvate is achieved by mixing ¹³C-labeled compounds with an electron paramagnetic agent (EPA), e.g., OXO-63. The mixed entity is then placed in a 3.35-T magnetic field, cooled to ~1 K, and microwaves are used to transfer polarization from the electron spin of the EPA to the ¹³C nuclei of the biomolecule [66]. Once the polarization is achieved, the sample is rapidly dissolved with hot, sterile water and neutralized to physiological pH, temperature, and osmolality. Hyperpolarized ¹³C pyruvate can be intravenously injected where it metabolized in tissues to various metabolites such as bicarbonate , lactate , and alanine depending on the dominant metabolic pathway in the tissue of interest in vivo. The conversion of pyruvate to each of these metabolites can be imaged using chemical shift MRI or MR spectroscopic imaging (MRSI) . The data acquisition utilizing hyperpolarized pyruvate is very rapid, approximately 60 s for [1^-13C] pyruvate at 3 T. Several studies have shown the preclinical application of using hyperpolarized pyruvate [64, 65]. Studies examined hyperpolarized ¹³C pyruvate and lactate in a C6 glioma-bearing animal before radiotherapy (Fig. 5.3b–d, top) and 96 h after radiotherapy (Fig. 5.3b–d, bottom). Comparison of spectra from tumor voxels with those on the contralateral side of the brain indicated high lactate signal in the tumor than in the brain (Fig. 5.3a). ¹³C chemical shift images acquired following intravenous injection of hyperpolarized [1^-13C] pyruvate into rats with implanted C6 gliomas showed significant labeling of lactate within the tumors but comparatively low levels in the surrounding brain. Labeled pyruvate signal was observed at high levels in blood vessels above the brain and from other major vessels elsewhere but was detected at only low levels in the tumor and in the brain [67, 68]. The ratio of hyperpolarized ¹³C label in tumor lactate compared with the maximum pyruvate signal in the blood vessels was decreased from 0.38 to 0.23 (a reduction of 34 %) by 72 h following whole-brain irradiation with 15 Gy (Fig. 5.3) [67]. Further studies in a transgenic adenocarcinoma of the mouse prostate (TRAMP) model demonstrated elevated levels of hyperpolarized [1^-13C] lactate in tumor, with the ratio of [1^-13C] lactate/[1^-13C] pyruvate being increased in high-grade tumors and decreased after successful treatment (Fig. 5.4) [69]. The preclinical studies using hyperpolarization of pyruvate has now advanced to human clinical trials. The first human translation of hyperpolarized technology was successfully demonstrated in patients with prostate tumor [70]. In brief, the study imaged 31 untreated patients diagnosed with localized prostate cancer, where 23 patients had Gleason score of 6, 6 patients with Gleason score 7, and 2 patients with Gleason score 8. The initial phase 1 of the study evaluated the safety, feasibility, and tolerability of injected hyperpolarized [1-¹³C] pyruvate. The study indicated no dose-limiting toxicities associated with hyperpolarized [1-¹³C] pyruvate. In addition, the median time taken for the dissolution of the agent to delivery into patient was 66 s. The ¹D dynamic MRSI data obtained showed higher [1-¹³C] lactate signal coming from tumor, and no detectable [1-¹³C] lactate signal came from the area which did not had tumor. The study confirmed correlation between the high lactate signals with prostate cancer grade. In conclusion, the study successfully demonstrated safety, tolerability, kinetics of hyperpolarized pyruvate delivery, and imaging hyperpolarized ¹³C metabolism (Fig. 5.5). In the future, more studies will be designed utilizing hyperpolarized [1-¹³C] pyruvate in larger cohorts of patients with different tumor types to firmly acknowledge the correlation with tumor grade and changes with therapy .

MRSI combines the ability of spectroscopy to acquire a large volume of metabolic information, with the ability of imaging to localize information spatially. Although phosphorus (³¹P) and carbon (¹³C) MRSI are possible, proton (¹H) MRSI is the technique most often used in clinical settings. On ¹H MRSI , tumor spectra contain resonances from metabolites such as taurine , total choline (choline, phosphocholine, and glycerophosphocholine), total creatine (phosphocreatine and creatine), pyruvate , and lactate .

Among the hallmarks of cancer described by Hanahan and Weinberg, elevated choline metabolism resulting in the accumulation of choline containing compounds such as choline , phosphocholine , and glycerophosphocholine (together represented as tCho) is recognized as an important hallmark, which is amenable for in vivo detection, by MRSI. Studies have shown that malignant transformation rather than just rapid proliferation as the reason for the increased accumulation of tCho is making this a very specific MRI-based imaging biomarker for proliferation [72, 73]. Tumor microenvironmental features such as hypoxia and acidotic extracellular conditions have also been associated with elevated tCho [74]. Several clinical studies already support the strength of monitoring tCho in detecting malignancies and predicting survival. Numerous multi-institutional clinical studies are examining the utility of MRS to detect tCho to aid in the diagnosis and treatment response monitoring .

Clearly, the ability to identify and quantify tumor oxygenation status and energy metabolism has far-reaching implications in a wide range of medical settings. Most importantly, a clinically useful assay to measure hypoxia , firstly, must be able to distinguish normoxic regions from the ones that are hypoxic at a level relevant to tumor oxygen partial pressure (pO₂) falling in the range of 5–15 mmHg. The ability to simultaneously image tumor oxygenation and metabolic profile can profoundly guide future therapies involving inclusion of radiosensitizers, hypoxia-directed cytotoxins, oxygen-enhanced gas mixtures such as carbogen (a mixture consisting of 95 % oxygen and 5 % carbon dioxide) [84], and hypoxia-activated prodrugs . Non-toxic prodrugs that generate active species in hypoxic tissue by selective bioreduction have now reached advanced clinical trials . Such hypoxia prodrugs mainly include tirapazamine [85], PR104 [86], and TH-302 [87]. In the early 1990s, hypoxia measurements were achieved by implanting properly calibrated, oxygen-sensitive electrodes (Eppendorf pO₂ histograph) which directly measured pO₂ in units of mmHg [88]. This technique had several practical disadvantages. It is limited to tumors that the probe can be easily accessed, and its invasive nature can cause tissue damage. However, this approach is still in use in the clinic. Other approaches such as immunohistochemistry of tumor tissue using extrinsic hypoxia-specific biomarker such as pimonidazole and intrinsic hypoxia biomarker such as carbonic anhydrase IX (CAIX) , VEGF receptor expression , and glucose transporter 1 (Glut1) are utilized to measure hypoxia [89]. These biomarker signatures show poor prognosis and are correlative to the extent of tumor hypoxia. However, the drawback of this approach is that it requires a biopsy, which is challenging and sometimes is not possible. Serum markers have also been evaluated to measure tumor hypoxia but with less success [90, 91]. Furthermore, sampled biomarkers are not able to evaluate spatial heterogeneity , which is often relevant to overall response and is essential for defining a radiation treatment field. Notably, all the above methods do not provide the longitudinal monitoring of hypoxia as the measurement is restricted to only a smaller sub-volume of a tumor. Noninvasive approaches have been instrumental in allowing serial imaging of hypoxia quantitatively and provide valuable information about cycling nature of hypoxia in both space and time. PET hypoxia imaging is noninvasive and routinely used in clinic . Due to the clinical significance of hypoxia imaging, an increasing number of hypoxia PET tracers are available and are being evaluated in the clinical setting. The first radionuclide detection of hypoxia in tumors was reported using ¹⁴C-misonidazole autoradiography [92]. Subsequently, two main tracer classes have been developed to specifically study hypoxia with PET. These tracers are ¹⁸F-labeled 2-nitroimidazole derivatives, such as [¹⁸F]-fluoromisonidazole ([¹⁸F]-FMISO) , [18F]-azomycinarabinoside [¹⁸F]-FAZA , [18F]-fluoroerythronitroimidazole [¹⁸F]-FETNIM , and Cu-labeled diacetyl-bis (N⁴-methylthiosemicarbazone) analogues. Among all these tracers , ¹⁸F-FMISO is the lead candidate and most extensively studied 2-nitroimidazole-based radiopharmaceutical PET tracer in the clinics. Clinical application of ¹⁸F-FMISO in radiotherapy planning and dose painting of different solid tumors is well described in Chap. 9. Despite its potential clinical application, ¹⁸F-FMISO imaging in rectal cancer was shown to be compromised by high nonspecific tracer accumulation in normoxic tissue [93]. Several other techniques to measure tumor hypoxia quantitatively and noninvasively are in preclinical development, namely, Overhauser MRI (OMRI), electron paramagnetic resonance imaging (EPRI) , and ¹⁹F MRI. Electron paramagnetic resonance (EPR) predominantly measures interstitial hypoxia. We describe here highly promising hypoxia imaging modalities that have gained immense attention in both preclinical and clinical setting .