Compared with optical and nuclear imaging methods, the lack of imaging sensitivity is a major limitation for MRI in molecular imaging applications (16
). Conventional small molecular T1
agents (e.g., Gd-DTPA) or new CEST probes [e.g., Eu-DOTA-4AmCE (17
)] have limited sensitivity of detection (>μmol/L), which makes it difficult to image specific tumor markers at low physiologic concentrations (<nmol/L). Superparamagnetic nanoparticles (e.g., Fe3
) have shown significantly improved sensitivity due to their strong perturbation to the local magnetic field. Currently, clinically used superparamagnetic agents are synthesized by aqueous precipitation of FeCl2
in the presence of a dextran polymer (e.g., Feridex). This method yields a variable size distribution of iron oxide nanoparticles (2–20 nm) and the loading of iron oxide per nanoparticle is low (18
). Compared with the Fe3
-dextran system, our SPPM design had several distinctive advantages. First, the size distribution of Fe3
nanoparticles was monodisperse (e.g., 9.9 ± 0.4 nm in diameter), which minimized variability between Fe3
nanoparticles. Second, clustering of highly compacted Fe3
particles inside micelle core considerably increased MR relaxivity. Previous study showed that over 10 times enhancement in T2
relaxivity per Fe was achieved when clustered Fe3
particles were loaded inside each micelle (10
). Third, high loading of Fe3
(e.g., 33 wt% in this study) effectively increased Fe content per micelle nanoparticle. Combination of increased molar relaxivity and high Fe loading per particle resulted in considerably increased sensitivity of SPPM probes. Indeed, results from this study showed detection of picomolar concentrations of SPPM nanoparticles in phantom samples by MRI. Moreover, the SPPM samples showed superb stability with long storage shelf-lives. These nanoparticles do not aggregate at 4°C even after 1 month of storage in PBS solution. DLS measurements also showed that particle size (~70 nm in diameter) did not change over this time period.
Pharmacokinetic studies showed that SPPM formulations had prolonged blood circulation times, which should allow for effective tumor targeting by the cRGD-encoded SPPM. Both cRGD-encoded and cRGD-free SPPM formulations had comparable α-phase plasma half-lives (t1/2,α
) at 0.34 ± 0.09 and 0.40 ± 0.34 hours, respectively. However, cRGD-free SPPM showed slower clearance in the β-phase as represented by longer t1/2,β
(9.2 ± 0.8 hours) than cRGD-encoded SPPM (3.9 ± 0.8 hours). We attribute this variation to the different functionalization of peptides (i.e., cRGD versus Cys) on the SPPM surface. Because the ORS imaging study was performed 1 hour after SPPM injection, we do not anticipate that the different t1/2,β
values will have a strong influence on the ORS contrast in the current study. Biodistribution studies showed relatively high accumulations of SPPM in the liver and spleen, which are commonly observed for nanoparticles in vivo
). In comparison, SPPM accumulations in other major organs, such as lung, brain, muscle, and kidney, were minimal. More importantly, tumor accumulation of cRGD-encoded SPPM was significantly higher than that of cRGD-free SPPM and cRGD-encoded SPPM coinjected with free cRGD ().
Currently, the T2*-w method is the gold standard for imaging superparamagnetic nanoparticles. Due to the strong magnetization and field perturbation by the superparamagnetic nanoparticles, the T2*-w method provides a much higher sensitivity for SPIO agents over conventional T1 agents (e.g., Gd-DTPA). Despite this advantage, T2*-w imaging suffers from several major limitations in molecular imaging applications. First, in T2*-w imaging, significant signal loss can arise due to B0 inhomogeneity, magnetic susceptibility, or spin-spin couplings. This will lead to signal variations that are independent of the superparamagnetic probes. For example, the preinjection T2*-w image () showed considerably different image contrast (tumor ROI analyses showed 2.1 ± 1.2 and 1.2 ± 0.9 for the SPPM-t and control-t, respectively) between the two tumor xenografts in the same mouse before SPPM injection. After SPPM injection, the control tumor (SPPM-free) showed a relative increase in signal intensity (1.7 ± 1.1) compared with its preinjection image (1.2 ± 0.9). Such signal variations can greatly complicate the interpretation of preinjection and postinjection images to identify SPPM contrast on an individual animal basis. Second, T2* method is also sensitive to different magnetic susceptibility caused by air/tissue or hard/soft tissue interface present in internal organs. These phenomena give rise to signal distortion and can also complicate image analysis. Lastly, identification of SPIO requires a precontrast scan for image subtraction from a postcontrast scan. Small changes of animal positions can easily decrease the spatial resolution (>mm) for tumor detection. Compounded with the variations in signal intensity between the scans, the quality and accuracy of contrast images can be considerably deteriorated in subtracted images.
The ORS method overcomes the above limitations and offers many advantages. The most exciting aspect of ORS is its ability to turn ON the contrast of the SPPM probes after the contrast agents are injected. This ability greatly increases the imaging accuracy in detecting contrast changes in targeted tissues while saving the need of a precontrast scan as in the T2*-w method. Because ON and OFF imaging can be performed subsequently without the moving of the imaging subject, pixel-by-pixel subtraction between these images can be performed to provide the maximal spatial resolution in contrast images. In this case, the intrinsic resolution of MRI (e.g., ~100 μm for the 4.7 T scanner in the current study) can be maintained in the ORS contrast images (e.g., and ). In contrast, it is impossible to perform accurate pixel-by-pixel subtractions between preinjection and postinjection images in T2*-w method. It is also interesting to note that although the current acquisition conditions with T2*-w and ORS methods allowed for similar sensitivities (e.g., S90) in phantom studies, results from animal studies showed significantly improved SNR/CNR by the ORS method for SPPM detection than the T2*-w method (; ). This improvement indicates that complex physiologic environments may introduce more signal variations in T2*-w images than the ORS images.
One consideration in using the ORS method is that the ORS imaging protocol also introduces contrast effects by magnetization transfer (MT). The MT effect is dependent on cross-relaxation and/or chemical exchange between the “free” water and macromolecule-associated or “immobile” water (20
), whereas ORS effect primarily relies on the diffusion of water molecules among different compartments that are defined by magnetic field isosurfaces (12
). Our current study allows for a preliminary estimation of MT and ORS effects on SPPM contrast. The CNR for the SPPM-injected tumor is 37.4 ± 8.4 (n
= 3), which includes the compounded effects from both ORS and MT. The CNR for the SPPM-free tumors from the same animal is 7.8 ± 6.8, where the SPPM-induced ORS contribution is absent. The significantly higher CNR (P
= 0.01) from SPPM-injected tumors suggests that ORS contrast can be effectively detected over the MT contrast. Precautions need to be taken when imaging lower concentrations of SPPM in tumor tissues when the contribution from MT effect becomes significant. Further studies are necessary to quantitatively evaluate the contributions of MT- and/or ORS-based effects on the detection of SPPM.
In conclusion, these studies show the synergy of ultrasensitive SPPM design and ORS method for molecular imaging of cancer, and in particular, non–small cell lung cancer. Phantom studies show detection limit at picomolar (10−12 mol/L) concentrations of SPPM nanoprobes, making them comparable in sensitivity with nuclear imaging probes. Results from the animal studies support our hypothesis that ORS method can significantly increase the contrast sensitivity and detection accuracy of SPPM particles in tumor tissues over the conventional T2*-w method. After i.v. injection, αvβ3-targeted SPPM nanoprobes show significantly increased ORS imaging contrast in A549 tumors over the nontargeted SPPM. Pharmacokinetic studies showed prolonged blood circulation times of SPPM nanoparticles and verified the αvβ3-dependent targeting specificity by the cRGD-encoded SPPM. The combination of ORS imaging with cancer-targeted SPPM nanoparticles offers new opportunities in detecting biochemical markers at early stages of tumor development.