Magnetic resonance susceptibility based perfusion imaging of tumors using iron oxide nanoparticles

Schematic illustrating: (a) random orientations of the magnetic domains in a superparamagnetic iron oxide (SPIO) particle in the absence of any applied magnetic field, and (b) application of an external magnetic field B₀ causing the magnetic domains of the SPIO particle to orient along B₀. (c) Superparamagnetic particles (S) (open arrows) have a very high magnetic susceptibility and can be magnetized (M) to saturation (M_S = saturation magnetization) even in weak external magnetic fields (H). Unmagnetized ferromagnetic and ferrimagnetic materials (F) (solid arrows) also become magnetized along this curve. However, once magnetized, ferromagnetic and ferrimagnetic materials retain their magnetization (M_R = remnant magnetization), even if the external field is reduced to zero. However, at ambient temperatures superparamagnetic materials do not retain their magnetization in the absence of an applied field. (Reprinted, with permission, from Ref. 25 Copyright 2001 Springer and Ref. 30 Copyright 1972 Addison‐Wesley).

(a) Schematic illustrating the three different regimes of susceptibility induced relaxation effects and the differential sensitivity of gradient‐echo (GE) () and spin‐echo (SE) (ΔR₂) relaxation rates to vessel caliber. This sensitivity to vessel size constitutes the basis for imaging macro‐ and microvascular blood volume as well as imaging vessel‐size. (b) Size dependence of and ΔR₂ for fractional volume = 2% and Δχ = 1 × 10⁻⁷. ΔR₂ peaks for microvessels, while for all radii and plateaus for macrovessels. (Reprinted, with permission, from Refs.14,21).

Schematic illustrating the origins of susceptibility based Magnetic resonance (MR) contrast. (a) In the absence of any susceptibility difference between blood (χ₁) and the surrounding tissue (χ₂), no microscopic magnetic field gradient is set up and diffusing water protons experience the same local magnetic field. (b) When a susceptibility difference (Δ χ) arises between the intravascular space and the surrounding tissue, say as a result of the presence of superparamagnetic iron oxide (SPIO) contrast agent, a microscopic field gradient (‐‐‐) is set up that perturbs the local magnetic field, and diffusing water protons experience different local magnetic fields, leading to loss of phase coherence, and MR signal attenuation that can be followed dynamically using either T₂‐ or ‐weighted MR pulse sequences. This constitutes the basis of dynamic susceptibility based contrast (DSC) magnetic resonance imaging (MRI). (c) Surface plot illustrating the three‐dimensional aspects of mathematically simulated microscopic magnetic field gradients induced around a microvessel. The orientation of the applied field (B₀) and the axis along which the normalized field change (Δ B/B₀ Δ χ) is plotted are shown in the inset. (Reprinted, with permission, from Ref.14. Copyright 2004 Elsevier).

(a) Post‐Gd MR image of a 9L gliosarcoma bearing rat brain illustrating the contralateral (C) and tumor (T) ROIs employed for MR relative cerebral blood volume (rCBV) histogram analyses. (b) rCBV map computed from the first‐pass of monocrystalline iron oxide nanoparticle (MION), overlaid on a high‐resolution anatomical image: note elevated rCBV in the tumor, a finding consistent with a larger fractional blood volume (FV) relative to the contralateral brain. (c) Post‐Gd MR image illustrating the contralateral (C) and tumor (T) grids that were sampled for stereological analyses. H&E stained tissue section of vessels perfused with Microfil (a silicone injection compound) at 20× magnification of (d) tumor tissue and (e) normal brain. (f) A histogram of the GE rCBV data showing that the tumor rCBV is shifted to the right with respect to the contralateral rCBV, a difference that is also apparent from (g) the histogram of the stereologically calculated FV. (Reprinted, with permission, from Ref. 35. Copyright 2000 Springer Verlag).

(a) Post‐Gd axial magnetic resonance imaging (MRI) of the rat brain illustrating the ROIs for the tumor (red) and contralateral (yellow) brain. (b) Post‐monocrystalline iron oxide nanoparticle (MION) axial MRI of the rat illustrating the ROIs for the gray (gray) and white (white) matter, respectively. Binarized images of tissue sections of microfilled vessels at 20× magnification of: (c) tumor, (d) contralateral brain, (e) gray matter, and (f) white matter ROIs. The color of each frame on the lower panel corresponds to the color of the ROI from which the tissue sections were sampled for histology. (g) Comparison of the ratios of fractional volumes obtained from MRI and histology for the tumor versus contralateral, and gray versus white ROIs, respectively. The error bars represent the standard error of the mean for each technique (N = 6 rats for the gray/white and N = 8 rats for the contra/tumor). There was no significant (P = 0.525) difference between gray/white and the ratio gray/white fractional volume but a significant (P = 0.005) difference between tumor/contra and the ratio tumor/contralateral fractional volume. This result indicates that the GE calibration factor is the same for gray and white matter but not the same for brain and tumor tissue. (Reprinted, with permission, from Ref. 39. Copyright 1994).

(a) Post‐Gd MR image of a 9L gliosarcoma bearing rat brain illustrating the tumor ROI (yellow hatched ellipse). (b) Post‐monocrystalline iron oxide nanoparticle (MION) ratio /ΔR₂ map of a 9L gliosarcoma bearing rat brain wherein one can clearly see the elevated ratio values in the angiogenic tumor rim (yellow hatched ellipse). (c) A histogram of the /ΔR₂ data showing that for the tumor ROI (over all slices), /ΔR₂ is shifted to the right i.e., larger caliber vessels with respect to the contralateral brain ROI (over all slices) /ΔR₂, a difference that is also apparent from (d) the histogram of the stereologically calculated vessel radii. (Reprinted, with permission, from Ref.62. Copyright 2001).