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Proc Natl Acad Sci U S A. Oct 20, 2009; 106(42): 17898–17903.
Published online Oct 6, 2009. doi:  10.1073/pnas.0908447106
PMCID: PMC2761243
Medical Sciences

Simultaneous imaging of tumor oxygenation and microvascular permeability using Overhauser enhanced MRI


Architectural and functional abnormalities of blood vessels are a common feature in tumors. A consequence of increased vascular permeability and concomitant aberrant blood flow is poor delivery of oxygen and drugs, which is associated with treatment resistance. In the present study, we describe a strategy to simultaneously visualize tissue oxygen concentration and microvascular permeability by using a hyperpolarized 1H-MRI, known as Overhauser enhanced MRI (OMRI), and an oxygen-sensitive contrast agent OX63. Substantial MRI signal enhancement was induced by dynamic nuclear polarization (DNP). The DNP achieved up to a 7,000% increase in MRI signal at an OX63 concentration of 1.5 mM compared with that under thermal equilibrium state. The extent of hyperpolarization is influenced mainly by the local concentration of OX63 and inversely by the tissue oxygen level. By collecting dynamic OMRI images at different hyperpolarization levels, local oxygen concentration and microvascular permeability of OX63 can be simultaneously determined. Application of this modality to murine tumors revealed that tumor regions with high vascular permeability were spatio-temporally coincident with hypoxia. Quantitative analysis of image data from individual animals showed an inverse correlation between tumor vascular leakage and median oxygen concentration. Immunohistochemical analyses of tumor tissues obtained from the same animals after OMRI experiments demonstrated that lack of integrity in tumor blood vessels was associated with increased tumor microvascular permeability. This dual imaging technique may be useful for the longitudinal assessment of changes in tumor vascular function and oxygenation in response to chemotherapy, radiotherapy, or antiangiogenic treatment.

Keywords: angiogenesis, dynamic nuclear polarization, hyperpolarized MRI, tumor hypoxia, DCE-MRI

Tumors >1–2 mm3 cannot grow further by relying on passive transport for nutrients and oxygen. To overcome this problem, tumors promote angiogenesis by releasing proangiogenic factors (1, 2). In normal tissue, there is a balance between proangiogenic and antiangiogenic factors, where mature vasculature with regular architecture ensures adequate supply of nutrients and oxygen. However, in tumors, because of excess proangiogenic factors, tumor neo-vasculature is abnormal and chaotic in architecture and has poor structural integrity (3). The lack of adequate pericyte and endothelial coverage results in large vascular pores causing a marked regional heterogeneity in tumor perfusion and making the tumor vascular hyperpermeable (4, 5). As a result, tumor progression is associated with disorganized angiogenesis, which results in inadequate oxygen supply and limited delivery of chemotherapeutics to the tumor (6, 7).

Noninvasive methods to obtain information pertaining to tumor microvessel density, vascular permeability, and oxygenation will aid in appropriate treatment choices (810). Monitoring the leakage of exogenously administered tracers from blood vessels can help in assessing tumor vascular permeability (4, 8) by dynamic contrast-enhanced MRI (DCE-MRI). These contrast agents diffuse from the blood circulation into the extravascular extracellular space (EES) at a rate determined by blood perfusion, vascular permeability, and surface area (11). The endothelial (volume) transfer coefficient Ktrans (12), can be obtained from pharmacokinetic analysis of the contrast agent-induced dynamic signal change that provides useful data to quantify the treatment response in anticancer therapy.

The interpretation of such kinetic information depends on the molecular size of the tracers used (12, 13). Small tracers (<1 kDa) such as 15O-labeled H2O in PET and gadolinium diethylenetriaminepenta-acetic acid (Gd-DTPA) in MRI can diffuse across the endothelial barrier of both normal (except for brain) and tumor tissues. Their blood-to-tissue transfer rate reflects the blood perfusion difference in tumor (11, 14, 15). However, the endothelial transfer of large contrast agents is distinctly sensitive to the microvascular permeability difference (13, 16). DCE-MRI techniques with large molecular mass tracers have been increasingly used in the assessment of tumor vascular permeability (8, 13, 15). However, despite the positive impact on patient survival, quantitative and noninvasive information regarding the relationship between tumor microvascular permeability and oxygen status is not available.

In this study, we have developed a noninvasive imaging modality, which can simultaneously obtain quantitative information pertaining to tumor oxygenation status (pO2) and microvascular permeability. The imaging modality is based on hyperpolarizing tissue water protons in vivo by using dynamic nuclear polarization (DNP) and an oxygen-sensitive paramagnetic agent OX63. OX63 (molecular mass 1427) has a molecular mass three times larger than typical gadolinium complexes and so its endothelial transport is more sensitive to microvascular permeability. Using Overhauser enhanced MRI (OMRI), the image intensity can be enhanced by hyperpolarizing the nuclear spin states of water protons surrounding OX63. The image intensification is directly proportional to the contrast agent concentration and inversely to the oxygen concentration (17, 18). From the dynamic profile of hyperpolarized enhancement of 1H-MRI signal after injection of OX63, it is possible to simultaneously determine the tissue oxygen concentration and the kinetic constants of OX63 related to tumor microvascular permeability.


DNP-Based Signal Enhancement in OMRI.

Fig. 1A shows spin-lattice relaxation time (T1)-weighted MR image of a five-tube phantom containing different concentrations of Gd-DTPA in the range of 0 to 1.5 mM obtained at field strength of 7 T. The image intensity of aqueous solutions was sufficiently high even in the absence of Gd-DTPA and increased in a Gd-DTPA concentration-dependent manner (Fig. 1B). Gd-DTPA (1.5 mM) results in a 2.4-fold signal enhancement. A similar phantom containing OX63 was prepared to study the image intensity enhancement profiles in OMRI at 0.015 T. As expected from the low field used for OMRI, in the absence of OX63 (0 mM, center tube), the image from the aqueous solution is of poor quality (Fig. 1C). However, in the presence of OX63 with hyperpolarizing radio frequency (RF) pulse, the image intensity substantially increased in proportion to OX63 concentration (Fig. 1D). Whereas image intensity enhancement was noticed both with Gd-DTPA in MRI at 7 T and OX63 in OMRI at 0.015 T, the mechanisms responsible for these enhancements are different. With gadolinium complex in T1-weighted MRI, the image intensity enhancement is achieved by shortening the T1 of water protons. However, in OMRI the image enhancement caused by OX63 results from induction of large hyperpolarization of water protons via a phenomenon known as DNP or the Overhauser effect (17). The dynamic range in terms of image enhancement is higher with OX63 in OMRI (0–7,000%; Fig. 1E, red circles) than with Gd-DTPA in T1-weighted MRI (100–300%; Fig. 1E, blue squares), providing superior quality in the parametric image (Fig. 1D).

Fig. 1.
DNP-based signal enhancement in OMRI. (A) T1-weighted 7-T MRI image of five tubes phantom containing 0-, 0.25-, 0.5-, 1.0-, and 1.5-mM solutions of Gd-DTPA. (B) Parametric signal enhancement map of MRI/Gd-DTPA calculated from A. (C) Hyperpolarized MRI ...

OMRI with OX63 Can Determine Tumor Vascular Functions and Oxygenation Status Simultaneously.

To evaluate the capability of OMRI to extract tumor vascular function and oxygenation status simultaneously, a squamous cell carcinoma (SCC) tumor bearing mouse (implanted on the right hind leg) was used. Fig. 2A shows the anatomic image of the tumor bearing mouse in an axial profile. The MRI image at 0.015 T without the hyperpolarizing RF pulse (Fig. 2B) after i.v. administration of OX63 had poor signal-to-noise ratio mainly caused by its low magnetic field. Application of the hyperpolarizing RF pulse resulted in a substantial enhancement of MRI image intensity in vivo (Fig. 2C). The extent of hyperpolarization, i.e., NMR signal enhancement, is influenced by the local concentration of hyperpolarizing agent and inversely by tissue oxygen level (17, 18). By varying the extent of hyperpolarization achieved by changing the intensity of the hyperpolarizing RF pulse, the local concentration of the OX63 probe (Fig. 2D) can be calculated from Eq. 3 (see SI Text for details). Although OMRI does not directly detect OX63, the OX63 molecule induces hyperpolarization of protons in adjacent water molecules, resulting in the intensification of 1H-MRI images. Taking into account the water diffusion distance <25 μm in the tumor tissue during the time between the hyperpolarizing pulse and image data collection ([similar, equals]50 ms), where the water diffusion constant is <0.002 mm2/s (9), the OMRI signal enhancement can be considered to originate from the location of the OX63 molecule. This process allows for the estimation of the endothelial transfer coefficient Ktrans by pharmacokinetic analysis of multitime OX63 concentration data from the dynamic OMRI experiments. A pixel-by-pixel calculation provided the Ktrans map from the image datasets (Fig. 2E), which were overlaid on an anatomic image obtained from conventional 7 T MRI. The tumor regions exhibited heterogeneous distribution of Ktrans values and had areas of relatively high Ktrans values (median Ktrans = 0.045 ± 0.011/min, n = 11, P < 0.01) compared with the muscle tissue region of the contralateral normal leg (0.020 ± 0.011/min, n = 11).

Fig. 2.
Simultaneous imaging of tissue pO2 and Ktrans of OX63 in vivo. (A) Axial anatomic image of SCC tumor-bearing mouse obtained by 7-T MRI. FOV = 32 mm. (B) Image after OX63 injection obtained by using OMRI scanner at 0.015 mT without hyperpolarizing RF irradiation. ...

The extent of hyperpolarization at a given concentration of the paramagnetic agent depends on the strength of the hyperpolarizing pulse and inversely on the local oxygen concentration (17, 18). This feature in OX63 makes it possible to noninvasively and quantitatively map the tissue pO2 by using Eq. 4 (details are available in SI Text), simultaneously with Ktrans maps using identical image datasets. The pO2 image was generated and coregistered with the anatomic image obtained at 7 T (Fig. 2F). In the pO2 image, substantial hypoxic regions can be seen in the SCC tumor. Quantitative examination revealed that ≈36% of the pixels in this tumor region had pO2 <10 mm Hg, whereas in the normal femoral muscle of the same animal ≈8.0% of the pixels had pO2 <10 mm Hg.

Comparison of the Two Modalities to Assess Tumor Vascular Permeability in the Same Mouse.

The Ktrans predominantly reflects blood perfusion, vascular permeability, and surface area, which is critically influenced by the molecular size of the contrast agents. Although both Gd-DTPA and OX63 are cell-impermeable contrast agents, the latter has approximately three times larger molecular mass. To compare the Ktrans values obtained from the two modalities, Gd-DTPA/T1-weighted MRI and OX63/OMRI experiments were conducted sequentially on the same tumor-bearing mouse. Fig. 3A shows the axial MRI image of the tumor-bearing mouse depicting morphological detail. The Ktrans map from a typical DCE-MRI experiment with Gd-DTPA at 7 T is shown in Fig. 3B. Relatively high Ktrans values were observed globally for Gd-DTPA compared with those for OX63 in both tumor and normal tissue regions, consistent with its lower molecular mass making it penetrate easily even through the smaller pores in normal blood vessels. Therefore, the Ktrans of Gd-DTPA may have larger contributions from blood perfusion in tissue than vascular permeability as reported (13, 15). OX63's larger molecular size makes its transfer through vascular endothelium in normal tissue more restricted relative to tumor vasculature. The Ktrans map of OX63 (Fig. 3C) was obtained from the same mouse from OMRI experiments. The Ktrans values of OX63 were lower in range (0–0.15 min−1) than those for Gd-DTPA (0–0.6 min−1) and displayed a marked heterogeneity in tumor regions. The KtransOX63 values were also lower in normal tissue compared with the tumor region, suggesting that KtransOX63 may be predominantly governed by the vascular permeability dependency of the OX63 probe rather than blood perfusion. Fig. 3D shows the pO2 maps derived from the same OMRI dataset used to compute the Ktrans of OX63. Fig. 3 E and F show magnified KtransOX63 and pO2 images from the tumor region selected from Fig. 3 C and D. Region of interest (ROI) 1 in Fig. 3E that shows the lower values of KtransOX63 is spatially coincident with relatively higher tumor pO2 values in Fig. 3F. Conversely, ROI-2, which has higher KtransOX63 values, is spatially coincident with regions exhibiting hypoxia. These observations are consistent with those made in prior studies that the hyperpermeable tumor vasculature is associated with insufficient blood flow and oxygen transport, resulting in tumor hypoxia (19, 20).

Fig. 3.
Comparison of Ktrans maps of Gd-DTPA and OX63 in a SCC tumor. (A) SCC tumor region can be distinguished in a T2-weighted anatomic image by using 7-T MRI. FOV = 32 mm. (B) Ktrans map of Gd-DTPA using 7-T MRI in SCC tumor-bearing mouse. (C) Ktrans map of ...

Tumor pO2 Inversely Correlate with the Ktrans of OX63.

The hyperpolarization-based MRI using OX63 allows for the comparison of tumor vascular function and oxygen level on the same animal by using the same image datasets. Fig. 4A shows the scatter plot of pO2 values in each pixel against Ktrans of OX63 obtained from the same tumor shown in Fig. 3. The pO2 declined with increase in the KtransOX63 in tumor regions. Inverse correlation was also observed between the median tumor pO2 and KtransOX63 obtained from 11 individual mice (Fig. 4B). If the main contribution to KtransOX63 is blood perfusion, positive relation would be expected between the pO2 and KtransOX63. The inverse relationship suggests that the leaky and abnormal tumor vasculature cannot adequately deliver oxygen into the tissue, in agreement with earlier studies using immunohistochemical analyses or oxygen sensitive electrodes (19, 20). It is noteworthy that the SCC tumors display minimal necrotic regions at this time point (<10 days after tumor implantation). The necrotic region may have very low KtransOX63 and pO2 values caused by limited blood perfusion. The existence of such necrotic regions might not exhibit the observed inverse relationship between pO2 and KtransOX63 in the tumor.

Fig. 4.
Correlation between tumor pO2 and KtransOX63. (A) Scatter plot of tumor pO2 values in each pixel against Ktrans of OX63 obtained from the same tumor shown in Fig. 3. (B) Simultaneous imaging of pO2 and KtransOX63 was performed in 11 different SCC tumor-bearing ...

Lack of Perivascular Cell Coverage May Be Responsible for Increased Tumor Microvascular Permeability.

Insufficient perivascular cell coverage has been attributed to the increased microvascular permeability in and around the tumor (20, 21). We investigated whether the tumor KtransOX63 values correlate with the immunostaining of α-smooth muscle actin (αSMA), a marker of pericytes. Fig. 5A shows the anatomic image of the SCC tumor obtained with 7 T MRI, and the corresponding KtransOX63 map by OMRI is shown in Fig. 5B. After MRI measurements, mice were killed and tumor tissues were carefully dissected so as to have the same slice as the one used to obtain the KtransOX63 image based on the T2-weighted anatomic MRI images, which in turn was confirmed by comparing with H&E staining results as shown in Fig. 5C. Immunostaining of αSMA expression revealed that tumor blood vessels in regions with high KtransOX63 values (regions 1 and 2 in Fig. 5B) were poorly covered by perivascular cells compared with the regions with limited endothelial transfer coefficient of the OX63 molecule (region 3 in Fig. 5C). A scatter plot of αSMA-positive area against corresponding local KtransOX63 values obtained from three to four different regions in each tumor tissue showed an inverse relation (Fig. 5F). These results suggest that lack of perivascular cell coverage in tumor blood vessels increases microvascular permeability and results in high KtransOX63 values and low pO2 (hypoxia).

Fig. 5.
Tumor KtransOX63 inversely correlates with vascular maturation. (A) Coronal T2-weighted 7-T MRI image of a SCC tumor-bearing mouse (mouse no. 1). FOV = 32 mm. (B) Corresponding KtransOX63 map obtained by OMRI. (C) H&E staining of whole tumor tissue ...


Normalization of tumor vasculature can improve the efficiency of both chemotherapy and radiotherapy. Tumor vascular maturation using methylselenocysteine improved tumor perfusion and the delivery of doxorubicin into tumors presumably by minimizing vascular permeability (21). Kashiwagi et al. (20) recently demonstrated that tumor vascular normalization by establishing a perivascular nitric oxide gradient decreased microvascular permeability and increased tumor oxygenation, resulting in improved response to radiation treatment. Tumor reoxygenation by antiangiogenesis drugs via vascular normalization and consequent radiosensitization were reported by Dings et al. (19) and Ansiaux et al. (22). However, the vascular normalization and resulting tumor reoxygenation are observed only in a narrow time period (19, 22), termed the “normalization window” by Jain et al. (7). Although several other oxygen imaging techniques, including 19F-MRI of perfluorocarbons and EPR imaging, have been reported, there are no known methods capable of simultaneously visualizing both tumor vascular functions and oxygenation status that allows for longitudinal monitoring before, during, and after treatment. In the present study, the use of the oxygen-dependent hyperpolarizing agent OX63 itself as a marker of microvascular permeability enabled the simultaneous imaging of tumor oxygen levels and microvascular permeability without spatial and temporal mismatching.

Quantifying the relationship between tumor oxygenation and vascular functions has a profound impact on preclinical and clinical cancer research. A positive correlation between pO2 from polarographic measurements and blood perfusion was shown in patients with head and neck SCC (23). Although tumor oxygenation status correlates with regional blood perfusion levels, the contribution of highly leaky neo-vasculature to tumor oxygen level remains unclear. In this study, tumor regions with high vascular permeability (Ktrans of OX63) were found to be hypoxic regardless of blood perfusion as determined by MRI using Gd-DTPA (Ktrans of Gd-DTPA). We also recently reported that significant hypoxia existed even in tumor regions that exhibited blood perfusion, whereas a positive correlation was observed between tumor oxygenation status and blood perfusion (9). These observations suggest that diffusion-limited hypoxia does not completely determine the oxygen status of tumors. Longitudinal oxygen gradient (24) and fluctuations in tumor blood flow (25) may exist and have consequences on delivery of oxygen and nutrients.

Hyperpolarization of nuclei in molecules by DNP is an emerging technique that substantially increases the sensitivity of MRI beyond the thermal equilibrium state. The increase in sensitivity has permitted MRI using hyperpolarized molecules as tracers. Hyperpolarized compounds labeled by 13C such as pyruvic acid and bicarbonate have been used to provide information related to metabolism and pH, marking a milestone in physiologic MRI (2628). With these methods, the hyperpolarization procedure was carried out in a polarizer outside of the MRI scanner at cryogenic temperatures (≈1 K). After hyperpolarization states are attained, the tracer is thawed quickly to room temperature while retaining its high nonequilibrium polarization and quickly administered to the living objects. Once administered, the spectral signatures of the original hyperpolarized tracer and the metabolites can be visualized (2628). However in OMRI, the hyperpolarization of the nuclear spin states of tissue water protons is accomplished in vivo after injecting the paramagnetic contrast agent by RF irradiation at the EPR frequency of the agent (18, 29). Although the penetration depth limitation of EPR irradiation pulse into biological tissue forces the use of a low magnetic field, the sensitivity disadvantage caused by this low magnetic field can be compensated with the intensity enhancement by DNP. The extent of hyperpolarization depends on the concentration and the EPR line width of OX63. The oxygen-dependent EPR line width of OX63 allows computation of tumor pO2, simultaneously with tumor vascular permeability of OX63 (see SI Text for details).

In conclusion, we have presented a noninvasive method to obtain in vivo microvascular permeability distribution simultaneously with quantitative pO2 maps. This process permits us to examine the relationship between tumor hypoxia and vascular functional integrity. The technique can be applied in experimental animals for repeated assessment of tumor vascular permeability and oxygen status without any detrimental effects. These features will be highly useful in assessing physiological changes in tumors in response to therapy that manifest earlier than tumor volume changes. Such imaging methods in small animals will be valuable in drug discovery research that has potential for human clinical use.

Materials and Methods


The triarylmethyl radical OX63 (17, 18) (obtained from GE Healthcare) has optimal chemical, pharmacological, and EPR characteristics such as stability, water solubility, low toxicity, long in vivo half-lives (17–21 min), single narrow line resonance, and pO2-dependent EPR line widths. The dose of OX63 used for OMRI (1.125 mmol/kg) was well below the maximally tolerated dose of 2.5–7.0 mmol/kg and the LD50 of 8 mmol/kg (18). Gd-DTPA was purchased from BioPAL. Other materials used were of analytical grade.

OMRI Scanner and Pulse Sequence.

A custom-built scanner (Philips Research Laboratories) operating in a field-cycled mode was used to avoid excess power deposition during the EPR cycle and improve sensitivity for NMR detection. Details on the OMRI scanner and pulse sequence are available in previous reports (18, 30). The OMRI experiments were performed by using a modified gradient echo sequence, where each phase-encoding step was preceded by an EPR saturation pulse to elicit the DNP. The pulse sequence started with the ramping of the B0 field to 8.1 mT. It was followed by the EPR irradiation pulse at 226 MHz, and ramping the B0 field to 15 mT, corresponding to NMR frequency of 625 kHz, before the NMR excitation pulse and the associated field gradients were turned on. For pO2 and Ktrans imaging, the OMRI images were collected with interleaving EPR irradiation pulses at two different power levels (90 and 11.3 W). Typical scan conditions are EPR irradiation time TEPR = 400 ms; echo time (TE) = 25 ms; repetition time (TR) = 800 ms; no. of average = 2; phase-encoding steps of 64. The pixel size was 0.75 × 0.75 mm2 with a slice thickness of 5 mm or 0.5 × 0.5 mm2 with 6-mm slice thickness. To obtain the pharmacokinetic profiles of the OX63, dynamic scans were performed for 15 min after injection, and at the end of the scans a conventional MRI (without EPR irradiation pulse) was collected for computing the enhancement factors. Image reconstruction and image processing to calculate quantitative parameters (pO2 and Ktrans) were done by using programs coded in MATLAB (MathWorks).

Computation of Tissue pO2 and Ktrans of OX63.

When a paramagnetic molecule is irradiated at the EPR frequency, the electron spin polarization is transferred to the surrounding water proton via DNP. The extent of enhancement in MRI signal intensity depends on the concentration (COX63) and intrinsic line width (a3 = 4.0 μT) of the OX63, the strength of the EPR RF power, and oxygen concentration (CO2) (18). The effect of these parameters on the line width of OX63 LWOX63 and resultant the transient magnetization of water proton M I are given by:

equation image

equation image

where γe is the gyromagnetic ratio of an unpaired electron spin, T1e and T2e are the electron spin relaxation times, α = 4.6 μT/W1/2 is the conversion efficiency of the EPR irradiation coil, P is the supplied RF power for EPR irradiation, and M01. is proton magnetization under thermal equilibrium state. Spectroscopic constants involving the oxygen-dependent and OX63 concentration-dependent line width change are a1 = 57 μT/mM O2 and a2 = 0.2 μT/mM OX63, respectively. If OMRI images are obtained at two different EPR power levels Pa and Pb, COX63 and CO2 can be estimated from:

equation image

equation image

where Einf* = Einf [1 − exp(−TEPR/T1] and Ia and Ib are the pixel intensities of high- and low-power images normalized with respect to the native MRI image. The T1 relaxivity of OX63 (r1) is 0.3 mM−1·s−1 (17), and the spin-lattice relaxation time of protons in the absence of OX63 T10 = 800 ms was measured and used for computation in the tumor region (30). A Patlak plot was used to graphically analyze the endothelial transfer constant Ktrans of OX63 from dynamic image datasets (31), which is expressed as:

equation image

where t is time after injection, Cp(t) is the plasma concentration of OX63, and V0 is the distribution volume of OX63 in the central compartment. Theoretical details of computation of tissue pO2 and Ktrans of OX63 are available in SI Text.

MRI Scanner and Pulse Sequences.

MRI measurements were done with a 7-T scanner controlled with ParaVision 3.0.2 (Bruker Bio Spin MRI). For comparison purposes, the Gd-DTPA phantom measurement (Fig. 1A) was designed to have the same spatial resolution to the OMRI image of the OX63 phantom. Axial T1-weighted fast low-angle shot (FLASH) images were obtained with TE of 3.5 ms, TR of 150 ms, flip angle of 45°, slice thickness of 5 mm, in-plane resolution of 0.75 × 0.75 mm2, and four averages.

In animal experiments, after a quick measurement of the sample position by a FLASH tripilot sequence, coronal and axial T2-weighted anatomical images were obtained by using a fast spin echo (RARE) sequence with TE of 13 ms, TR of 2,500 ms, 10 slices, RARE factor 8, resolution of 0.125 × 0.125 mm2, slice thickness of 2 mm, and acquisition time of 80 s. For T1 mapping, axial RARE images of three slices passing through the tumor region were obtained with TR values of 300, 2,000, and 3,000 ms. FLASH sequence of the same three slices was applied for DCE-MRI study of Gd-DTPA. The scan parameters are as follows: TE = 3.25 ms, TR = 150 ms, flip angle 45°, three slices, 0.375 × 0.375-mm2 resolution, 19.2-s acquisition time per image, and 100 times repetition.

In Vivo Experimental Design.

Female C3H/Hen mice were supplied at 6 weeks of age by the Frederick Cancer Research Center, housed five per cage in a climate-controlled, circadian rhythm-adjusted room, and they were allowed food and water ad libitum. Murine SCC VII tumor cells (5 × 105 viable cells) were injected into the s.c. space of the right hind leg of mice 7–10 days before the start of OMRI measurements. The tumor size during experimentation was ≈1.0–1.2 cm as measured by caliper. Mice were anesthetized by isofluran and fixed prone on a special holder. The tail vein was cannulated for administration of OX63 (1.125 mmol/kg body weight) and Gd-DTPA (0.125 mmol/kg body weight) solutions. After OMRI scans, mice were transferred into a 7-T MRI scanner without being removed from the holder to obtain the T2-weighted anatomic image. For Ktrans comparison purposes, three mice were also applied to DCE-MRI study upon Gd-DTPA injection. All experiments were carried out under the aegis of a protocol approved by the National Cancer Institute Animal Care and Use Committee and were in compliance with the Guide for the Care and Use of Laboratory Animal Resource, National Research Council.

Immunostaining of Pericyte Marker αSMA for Vascular Maturation Analysis.

After OMRI and 7-T MRI experiments, tumor tissues (≈5-mm thickness) corresponding to the OMRI image slice were exsected from mice at the same tilt angle and depth from the surface coinciding with the anatomic 7-T MRI images. OMRI image obtained with hyperpolarizing pulse has anatomic depiction and can be used for superimposing it to more detailed anatomic image by 7-T MRI, so that ROIs selected on a 7-T MRI image can be transferred to a KtransOX63 map by OMRI. Tumor tissues were fixed with 4% paraformaldehyde, frozen using ultracold ethanol, and sectioned 10 μm thick. After blocking nonspecific binding sites, the slides were left in the appropriate dilution of αSMA antibody (1:500; Abcam). The sections were incubated with Alexa Fluor-546 anti-rabbit secondary antibody (Molecular Probes; 1:500). Then they were mounted with Prolong Gold antifade reagent containing DAPI (Molecular Probes). ROIs of αSMA counts on the tissue section were decided under optical microscope mode in reference to anatomic image by 7-T MRI. Then the light switch was changed to fluorescent mode to obtain the fluorescent image of the selected locations. Serial sections were also stained with hematoxylin and eosin. The quantification of αSMA expression was done according to the method described by Zhou et al. (32).


Differences in KtransOX63 values of tumor and normal muscle regions in the same animal were compared by using the paired two-tailed Student's t test. Differences were considered significant when P < 0.05.

Supplementary Material

Supporting Information:


We thank the National Institutes of Health Fellows Editorial Board for editing this manuscript. This work was supported by the Intramural Research Program, Center for Cancer Research, National Cancer Institute, National Institutes of Health.


The authors declare no conflict of interest.

This article contains supporting information online at www.pnas.org/cgi/content/full/0908447106/DCSupplemental.


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