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Cell surface receptor-targeted magnetic iron oxide (IO) nanoparticles provide molecular magnetic resonance imaging (MRI) contrast agents for improving specificity of the detection of human cancer.
The present study reports the development of a novel targeted IO nanoparticle using a recombinant peptide containing the amino-terminal fragment (ATF) of urokinase plasminogen activator conjugated to IO nanoparticles (ATF-IO). This nanoparticle targets urokinase plasminogen activator receptor (uPAR), which is overexpressed in breast cancer tissues.
ATF-IO nanoparticles are able to specifically bind to and be internalized by uPAR-expressing tumor cells. Systemic delivery of ATF-IO nanoparticles into mice bearing subcutaneous and intraperitoneal mammary tumors leads to the accumulation of the particles in tumors, generating a strong MRI contrast detectable by a clinical MRI scanner at a field strength of 3 Tesla. Target specificity of ATF-IO nanoparticles demonstrated by in vivo MRI is further confirmed by near infrared (NIR) fluorescence imaging of the mammary tumors using NIR dye-labeled ATF peptides conjugated to IO nanoparticles. Furthermore, mice administered ATF-IO nanoparticles exhibit lower uptake of the particles in the liver and spleen compared to those receiving non-targeted IO nanoparticles.
Our results suggest that uPAR-targeted ATF-IO nanoparticles have potential as molecularly-targeted, dual modality imaging agents for in vivo imaging of breast cancer.
Breast cancer is the most common type of cancer and the second leading cause of cancer-related death among women. Novel approaches for the detection of primary and metastatic breast cancers are urgently needed to increase the survival of patients. A promising strategy to improve the specificity and sensitivity of cancer imaging is to use biomarker target-specific imaging probes (1–3) for image-based diagnosis and treatment monitoring. Currently, targeted radionuclide probes have been used for cancer detection by positron emission tomography or single photon emission tomography (2, 4, 5). Although nuclear imaging modalities show a high sensitivity, they lack good resolution and anatomic localization of the tumor lesion and require complicated and expensive radiochemistry. In addition, the half-life of the radiotracer often limits the ability for dynamic and time-resolved imaging and may not be able to capture the biomarker targeting agent to reach and accumulate in the tumor. Magnetic resonance imaging (MRI) offers a high spatial resolution and three-dimensional anatomic details and has been widely used in clinical oncology imaging. Recently breast MRI was recommended by the American Cancer Society as a screening approach, adjunct to mammography, for the early detection of breast cancer in women at high risk of this disease (6). Although breast cancer MRI shows a high sensitivity in detecting small breast lesions, a major challenge is its low specificity when using a non-targeted contrast agent such as gadolinium chelates, resulting in a high false positive rate and unnecessary biopsy and mastectomy (7). Therefore, the development of molecularly-targeted MRI contrast agents may increase the specificity of MRI as well as provide information on the level of biomarker expression in breast cancers.
Recently, biocompatible and functionalized nanoparticles have been shown to target tumors and produce optical, magnetic and/or radioactive signals for enhancing sensitivity and specificity of non-invasive tumor imaging. Previous studies have shown the feasibility of producing such imaging probes for in vivo MRI of cancers (8–10). The unique features of nanoparticles that make them suitable for receptor-targeted imaging include: 1) having a prolonged blood retention time, and 2) providing reactive function groups and a large surface area for loading large numbers or multiple types of tumor targeting ligands. However, several issues remain to be addressed in making nanoparticle imaging probes for clinical applications. These include the identification and availability of suitable imaging biomarkers, the delivery of sufficient levels of probe in vivo, and the development of imaging probes with sufficient signal amplification and contrast enhancement (11).
Paramagnetic iron oxide (IO) nanoparticles can induce remarkably strong MRI contrast, and have a large surface area and versatile surface chemistry for surface functionalization and the introduction of biomolecules (12–14). Their biological safety in humans has been tested, with non-targeted IO nanoparticles currently in use to detect liver tumor lesions or lymph node metastases in patients (15, 16). Previous attempts to develop targeted IO nanoparticles used dextran or poly(ethylene glycol) (PEG)-coated IO to conjugate targeting ligands and demonstrated the feasibility and improved sensitivity and specificity of such biomarker-targeted MRI nanoprobes (3, 14, 17–19). However, most of these ligands were directed to molecular targets that were expressed only in a small percentage of tumor tissues or subpopulations of tumor cells, such as Her-2/Neu, transferrin, MUC-1 and folate acid, which limits the sensitivity and application for molecularly-targeted cancer imaging in patients.
In this study, we used the amino-terminal fragment (ATF) of the high affinity receptor binding domain of urokinase plasminogen activator (uPA) to target its cellular receptor (uPAR), which is up-regulated in a high percentage of tumor cells and tumor-associated stromal cells, such as endothelial cells, macrophages and fibroblasts, of many human cancer types (20–23). It is well known that the interaction of uPA with its cellular receptor (uPAR) results in conversion of plasminogen to serine protease, a central regulator of the activation of other proteases including matrix metalloproteinase, which promotes tumor metastasis and angiogenesis. An elevated level of uPAR is associated with tumor aggressiveness, the presence of distant metastasis, and poor prognosis in breast cancer patients (24–28). In human breast cancer tissues, high levels of uPAR are detected in 54% of ductal carcinoma in situ (DCIS) and in 60 to 90% of invasive breast cancer tissues (24, 29).
To achieve optimal tumor targeting and imaging, we have developed novel paramagnetic IO nanoparticles that have uniform core sizes and are functionalized through surface coating of amphiphilic polymers. This surface coating provides a stable hydrophobic protective inner layer around a single crystal of IO nanoparticle with carboxylate groups in the outer layer readily available for conjugation with ATF peptides. Conjugated ATF-IO nanoparticles specifically bind to and are internalized by uPAR-expressing cells. We showed that systemic delivery of ATF-IO nanoparticles into mice bearing mammary tumors led to the accumulation of the nanoparticles in subcutaneous and metastatic mammary tumors, inducing MRI contrast changes sufficient for in vivo tumor imaging. By conjugation of Cy5.5, a near infrared (NIR) dye, to the ATF peptides, we were able to achieve in vivo tumor imaging with both MRI and optical imaging.
Mouse mammary carcinoma cell line 4T1 stably expressing a firefly luciferase gene was obtained from Dr. Mark W. Dewhirst at Duke University (Durham, NC). Human breast cancer cell line T47D was purchased from American Type Culture Collection (ATCC, Rockville, MD).
A cDNA fragment encoding amino acids 1−135 of mouse uPA, isolated by PCR amplification using a PCR primer pair containing forward (5′-CACCATGGGCAGTGTACTTGGAGCTCC-3′) and reverse (5′-GCTAAGAGAGCAGTCA-3′) primers, was cloned into pET101/D-TOPO expression vector (Invitrogen, Carlsbad, CA). Recombinant ATF peptides were expressed in E. coli BL21 (Invitrogen) and purified from bacterial extracts under native conditions using a Ni2+ NTA-agarose column (Qiagen, Valencia, CA). Purification efficiency was determined by SDS-PAGE and greater than 95% of purified proteins were ATF peptides. Cy™5.5 maleimide (GE Healthcare, Piscataway, NJ), a near infrared dye, was conjugated to ATF peptides using the manufacture’s protocol. Non-targeted dye molecules were separated from the Cy5.5 dye labeled ATF peptides using Sephadex G25 column.
Paramagnetic IO nanoparticles were prepared as described (30) using iron oxide powder as the iron precursor, oleic acid as the ligand and octadecene as the solvent. The core size and hydrodynamic size of the IO nanoparticles were measured using transmission electron microscopy (TEM), and light scattering scan, respectively. The particles were coated with amphiphilic polymers using a similar method as reported previously (31). ATF peptides were conjugated to the surface of IO nanoparticles via cross-linking of carboxyl groups to amino side groups on the ATF peptides as shown in Figure 1. Briefly, the polymer-coated IO nanoparticles were activated with ethyl-3-dimethyl amino propyl carbodiimide (EDAC, Pierce, Rockford, IL) and sulfo-NHS for 15 min. After purification using Nanosep 100k OMEGA (Pall Corp, Ann Arbor, MI), activated IO nanoparticles were reacted with ATF or Cy5.5-ATF peptides at a molar ratio IO:ATF of 1:20 in pH 7.0 PBS buffer at 4°C overnight, generating ATF-IO or Cy5.5-ATF-IO nanoparticles. The final ATF-IO conjugates were purified using Nanosep 100k column filtration. Conjugation efficiency of ATF peptides to the IO nanoparticles was confirmed by the measurement of fluorescence intensity of Cy5.5-ATF-IO nanoparticles using fluorescence spectroscopy and the ζ potential change before and after conjugation of the ATF ligands to IO nanoparticles. The number of ATF peptides conjugated to each IO nanoparticle was estimated by measuring the fluorescence intensity of a diluted sample of Cy5.5-ATF-IO nanoparticles using an emission wavelength of 696 nm and then comparing the value to a linear standard curve prepared using various concentrations of Cy5.5-ATF peptides.
Western blot analysis was performed as described using a standard protocol in our laboratory (32). To confirm the presence of ATF peptides in SDS/PAGE gel, the protein was transferred to PVDF membranes (Bio-Rad laboratories, Hercules, CA). An anti-His tag monoclonal antibody (Novagen, Madison, WI) was used to identify the His-tagged ATF-peptides. After reacting with HRP-labeled rat anti-mouse IgG antibody, the ATF peptide band was detected by enhanced chemiluminescence using ECL Plus (GE Healthcare) followed by autoradiography. The level of uPAR was determined using an anti-uPAR polyclonal rabbit antibody (Santa Cruz Biotechnology, Santa Cruz, CA) that reacts with both mouse and human uPAR and HRP conjugated goat anti-rabbit IgG. The protein bands were detected using ECL.
To determine whether the purified recombinant ATF peptides, either free or conjugated to IO nanoparticles, bind to uPAR, we performed a combined pull-down and Western blot analysis. Ni2+-NTA-agarose beads were incubated with His-tagged ATF peptides or ATF-IO nanoparticles at 4ºC for 30 min. The conjugated beads were washed twice with binding buffer and incubated with 500 µg of total cell lysate obtained from 4T1 or T47D cells for 2 hrs. Bound proteins were eluted from the beads using elution buffer containing 400 mM imidazole and then examined by Western Blot analysis to determine the amount of uPAR pulled down by ATF- or ATF-IO conjugated Ni-NTA agarose beads in each sample as described above.
Acetone-fixed frozen normal breast and cancer tissue sections were incubated with 5 µg/ml of polyclonal rabbit anti-uPAR antibody followed by biotinylated-goat anti-rabbit IgG. After further incubation with Texas-red avidin, the slides were examined under a fluorescence microscope (Zeiss Axioplan with Axiovision software, Carl Zeiss MicroImaging, Inc, Thornwood, NY). To detect the level of uPAR expression in living cancer cells, cells were incubated with an anti-uPAR antibody at 4 °C for 30 min. After incubation with FITC-goat anti-rabbit IgG, cells were examined under the fluorescence microscope.
Cells cultured on glass chamber slides (Nalge Nunc International, Naperville, IL) were incubated with 13.5 pmol of Cy 5.5-ATF-IO or non-targeted IO nanoparticles at 37 °C for 3 hrs. After washing with the PBS, the slides were examined under a confocal microscope (Perkin Elmer Ultraview ERS, PerkinElmer Life and Analytical Sciences, Inc, Wellesley, MA). To localize the IO nanoparticles, the cells incubated with Cy5.5-ATF-IO or IO nanoparticles were fixed with 4% formaldehyde in PBS and stained with Prussian blue staining by incubating the cells with a 1:1 mixture of 5% potassium ferrocyanide and 5% HCl acid for 30 min at 37 °C to confirm the presence of IO nanoparticles (33).
Cells were incubated with serum free medium containing non-targeted IO or ATF-IO nanoparticles at 37°C for 3 hrs. After washing with PBS buffer, the cells were embedded in 0.8% agarose in 24-well plates. Plates were then scanned in a 3T MRI scanner (Philips Healthcare, The Netherlands) using a T1-weighted gradient echo sequence and a multi-echo T2 weighted fast spin echo sequence which simultaneously collects a series of data points at different echo times (i.e., 20 TE points from 10–200 ms with 10 ms interval) for T2 relaxometry measurement.
Subcutaneous tumor model: Mouse mammary tumor 4T1 cells were injected subcutaneously into the back flank area of 6- to 8-week old female Balb/c or nude mice. We used nude mice for optical imaging to reduce background fluorescence.
Intraperitoneal (i.p.) metastatic mammary tumor model: 4T1 cells stably transfected with a firefly luciferase gene were directly injected into the upper right side of the peritoneal cavity. Tumor growth was monitored by bioluminescence imaging using the Xenogen bioluminescence imaging (BLI) system (Xenogen Corp., Hopkinton, MA).
Tumor-bearing mice were examined using a 3 Tesla MRI scanner (Philips Healthcare, The Netherlands), with a customized rodent coil to obtain MR images. Since MRI contrast effect is greatly dependent on the magnetic field strength (34), we chose to test the feasibility of detecting IO nanoparticle-induced contrast at a clinically relevant field strength (3T) with consideration of the potential application of this imaging probe in patients. Animals were anesthetized by i.p. injection of a ketamine:xylazine mixture (95:5 mg/kg). Animals were kept warm in the scanner using warm pads. A set of survey images was obtained using T2 weighted fast spin echo imaging sequence with TR of 5000 ms and TE of 20 ms. This was followed by high resolution imaging in the coronal view (i.e., slices cut through from head to tail) with a field of view (FOV) 110 × 40 mm, imaging matrix of 256 × 192 (reconstructed to 356 × 256), 40 slices with 1.1 mm slice thickness without slice gap.
The imaging sequences included T1 and T2 weighted spin echo or gradient echo methods, and the three-dimensional fast-spoiled gradient echo technique. A TE of 10 ms and TR of 350 ms were used for T1 weighted spin echo imaging, and TE of 50 ms, TR of 1500 ms for T2 weighted fast spin echo imaging. A multi-echo T2 weighted fast spin echo sequence with 12 TEs (range from 10–120 ms, 10 ms interval) was used to obtain T2 relaxometry of the whole mouse. The mice were injected with various nanoparticles suspended in PBS buffer though the tail vein and then scanned at different time points. For post-contrast scan, care was taken to maintain the same animal positions and same imaging sequences and parameters across different scan sessions. Images from pre- and post-contrast administration were compared to evaluate the efficacy of contrast enhancement. Multi-echo T2 images of slices were used for calculating T2 maps using a home-developed program based on Matlab (The Mathworks, Inc, Natick, MA). Because IO nanoparticles typically induced T2 weighted contrast change, the signal reduction in T2 weighted imaging and change in the T2 value were used to follow and estimate the accumulation of non-targeted or targeted IO nanoparticles in the area. Region of interest (ROI) analysis was used to evaluate and quantify the contrast agent-induced changes in MRI signal or T2 value in the tumor and other selected tissues and organs. ROIs of tumors were drawn based on the tumor T2 signal enhancement in T2 weighted images, or at a long TE time point (TE= 80 ms) in multi-TE T2 mapping imaging.
Tumor bearing mice were placed on an alfalfa-free rodent diet (Teklad #2918, Harlan Teklad, Madison, WI) to reduce background fluorescence. NIR images of the tumor-bearing mice were taken using the Kodak in vivo FX imaging system (Carestream Molecular Imaging, New Haven, CT) before and at different time points following the nanoparticle injection. For each NIR image, a corresponding X-ray image was taken to provide anatomic registration of the tumor.
Tumor and normal tissues were collected from the mice at the end of in vivo imaging experiments. 5 µm frozen tissue sections were incubated with Prussian blue staining solution to confirm the presence of IO nanoparticles in the tissue sections.
The magnetic IO nanoparticles synthesized for this study have a 10 nm core size. We coated the nanoparticles with a monolayer of amphiphilic polymers grafted with carbon alkyl side chains to stabilize and functionalize their surface, resulting in an 18 nm water soluble IO nanoparticle with carboxyl side groups. To reduce nonspecific binding and uptake by normal tissues, short PEG chains were conjugated to a portion of the carboxyl side groups on the amphiphilic polymers. Our results showed that magnetic IO nanocrystal used in this study has a strong T1 and T2 shortening effect with R1 (e.g., 1/T1) = 3.6 ± 0.3 mM−1.s−1 and R2 (1/T2) = 124 ± 7.2 mM−1.s−1 at 3T (Figure 1). Such amphiphilic polymer-coated IO nanoparticles have the combined characteristics of a relatively small particle size for easy in vivo delivery, a large surface area for conjugating biomolecules, and sufficient T2 effect for MRI contrast.
The amount and purity of recombinant mouse ATF peptides were determined by gel electrophoresis and Western blot analysis. Coomassie blue staining of the SDS-PAGE gel revealed an ATF band located at approximately 17 kDa (Figure 1). The presence of His-tagged ATF peptides was confirmed by Western blot using a monoclonal anti-His-tag antibody, showing a strong positive band in the location corresponding to the ATF-peptides identified by Coomassie blue staining (Figure 1). Conjugation of ATF peptides to the negatively charged carboxyl groups on the particle surface was further confirmed by the reduction of surface potential from an average of −30 mV to −11 mV after attachment of ATF peptides, suggesting the charge neutralization of carboxyl groups after conjugation with peptides (Figure 1). From the measurement of fluorescence intensity produced from the Cy5.5 dye conjugated to the ATF-IO nanoparticles, we estimated that 8 to 10 ATF peptides were attached to each nanoparticle (Figure 1).
We examined the level of uPAR expression in human breast cancer and normal tissues. A high level of surface uPAR was found in mouse mammary tumor 4T1 cells by immunofluorescence labeling of viable cells using an anti-uPAR antibody. On the other hand, human breast cancer T47D cells that lack uPAR expression were used as a negative control in this study (Figure 2A). Consistent with previous observations, uPAR is strongly expressed in invasive breast cancer tissues but is not detected in normal breast tissues (24) (Figure 2A)..
To determine whether ATF peptides maintain a high binding affinity after being conjugated to the IO nanoparticles, we performed a pull-down assay using cell lysates obtained from uPAR-positive 4T1 and -negative T47D cells. Our results showed that both free ATF peptides and ATF-IO nanoparticles bound to and precipitated uPAR proteins in the cell lysates, resulting in a positive uPAR band detected by Western blot assay in 4T1 but not T47D cell lysates (Figure 2B). It has been shown that interaction of uPA with uPAR leads to the internalization of the ligand/receptor complex (35). To determine whether the binding of ATF to uPAR leads to receptor-mediated endocytosis, we examined 4T1 and T47D cells after incubation with Cy5.5-ATF-IO at 37°C for 3 hrs under a confocal microscope. We found that Cy5.5-ATF-IO nanoparticles were internalized by 4T1 cells but not by T47D cells (Figure 2C). The presence of intracellular IO nanoparticles in 4T1 cells was further demonstrated by Prussian blue staining (Figure 2C).
Receptor-mediated binding and internalization of imaging probes promote their accumulation at the tumor site, which increases contrast effect. Prussian blue staining detected a high level of IO particles in 4T1 cells incubated with ATF-IO but not with non-targeted IO nanoparticles. uPAR-negative T47D cells showed only a very low level of nonspecific uptake (Figure 3A). 4T1 cells incubated with ATF-IO nanoparticles showed MRI contrast with signal reduction in T2-weighted gradient echo imaging (Figure 3B upper panel). T2 relaxometry measurements indicated that the T2 value of 4T1 cells bound with ATF-IO nanoparticles dropped significantly compared with those of the T47D cells incubated with ATF-IO nanoparticles and the samples treated with non-targeted IO nanoparticles (Figure 3B lower panel). Since the T2 value is a function of the iron concentration, the T2 relaxometry data suggest that reduction of T2 relaxation time and MRI signal in T2 weighted imaging are induced by the specific binding of ATF-IO nanoparticles to the uPAR expressing 4T1 cells.
In Balb/c mice bearing subcutaneous mouse mammary tumors derived from the 4T1 tumor cell line, T2-weighted gradient echo imaging and multiecho fast spin echo imaging showed that ATF-IO nanoparticles selectively accumulated in tumors, as evidenced by a reduction in T2 values and signal decrease in T2 weighted images in various areas of the tumor mass (Figure 4A). The ROI analysis of MRI signal change showed a three-fold signal reduction in animals receiving ATF-IO nanoparticles when compared with that in mice receiving non-targeted IO particles (Figure 4B). Although we observed decreases in MRI signals in the liver and spleen in ATF-IO-injected mice due to an IO particle-induced T2 effect, the reduction in MRI signal was 50% (liver) to 80% (spleen) less than that in mice that received non-targeted IO nanoparticles, suggesting that liver and spleen uptake of the nanoparticles was reduced for ATF-IO nanoparticles (Figure 4B). To further confirm the distribution of non-targeted and ATF-IO nanoparticles in normal and tumor tissues, Prussian blue staining was performed on tissue sections obtained from the mice that received control IO or ATF-IO nanoparticles. Prussian blue-stained cells were detected in the tumor sections of animals that received ATF-IO nanoparticles but not in sections from animals that received non-targeted IO nanoparticles. High magnification images showed the intracellular localization of IO nanoparticles in the cells (Figure 4C). In normal tissues, we found high levels of Prussian blue-positive cells in the liver and spleen of the mice that received non-targeted IO nanoparticles. However, liver and spleen tissue sections from the group receiving ATF-IO nanoparticles had fewer positive cells. We did not detect IO nanoparticles in the tissue sections of the brain or heart from mice injected with either non-targeted IO or ATF-IO nanoparticles (Figure 4C). For both groups, the lung and kidney tissues were also negative in most cases (Figure 4C) and only a few scattered iron-positive cells were detected in some sections.
We tested the feasibility of targeting and in vivo imaging of metastatic lesions using an animal model bearing i.p. 4T1 tumors. The presence and development of the 4T1 tumor was determined and followed by bioluminescence imaging (BLI) (36). At 5 hrs after injection of the ATF-IO nanoparticles, MRI signals decreased in two tumor lesions on top of the right kidney (Figure 5, upper panel). The MRI signal in the region recovered gradually 30 hrs after the injection of the ATF-IO nanoparticles. In contrast, we did not detect an IO nanoparticle-induced MRI signal change in a tumor-bearing mouse after injection of non-targeted IO (Figure 5, lower panel). The selective accumulation of uPAR targeted ATF-IO nanoparticles in this metastatic i.p. tumor model was further confirmed by Prussian blue staining of tissue sections from the sacrificed animals. A high percentage of iron-positive cells was found in the tumor lesion while the kidney beneath the tumor was negative (Figure 5, upper panel). On the other hand, we did not detect iron-positive cells in tissue sections of the i.p. tumor mass obtained from the mouse that received non-targeted IO nanoparticles, while the adjacent normal liver tissue had a high level of iron-positive cells (Figure 5, lower panel).
NIR optical imaging provides a simple and sensitive approach to confirm tumor targeting and to monitor the distribution of Cy5.5-ATF-IO nanoparticles in small animals. When monitoring mice bearing s.c. 4T1 tumors at different time points after injection of Cy5.5-ATF-IO nanoparticles, we observed a strong NIR signal in the tumor mass 24 hrs after the administration of nanoparticles. The signal intensity in the tumor mass increased to a peak level at 48 hrs but declined at 72 hrs (Figure 6A). We also detected NIR signals in kidneys, suggesting that some Cy5.5-ATF-IO nanoparticle conjugates and/or Cy5.5-ATF dissociated from IO nanoparticles may be eliminated through the kidney. Furthermore, we found that areas with MRI contrast change in the tumor overlapped well with the presence of NIR signal when comparing optical and MR images (Figure 6B). Examination of the tissue sections with positive Prussian blue staining revealed that Cy5.5 NIR signals co-localized with blue iron-positive cells (Figure 6C).
Previous studies reported that IO nanoparticles can be non-specifically taken up by macrophages as well as the reticuloendothelial system in the liver and spleen. We examined the tumor tissue sections stained with both Prussian blue and an antibody to CD68, a marker for macrophages (37). We found iron-positive cells present in both CD68-positive and -negative cell populations. A high percentage of iron-positive cells did not express CD68, suggesting that non-macrophage cell types, such as tumor cells, contain IO nanoparticles. It has been shown that active macrophages in breast cancer tissues also express a high level of uPAR (21, 38). It is possible that the presence of iron-positive macrophages may be due to active targeting of uPAR-expressing macrophages rather than non-specific uptake of IO nanoparticles.
Additionally, an MRI relaxometry T2 map of a mammary tumor showed that accumulation of ATF-IO nanoparticles was not uniformly distributed inside the tumor mass. At various levels of tumor MR images, MRI signal decreases were heterogeneous in the tumor, suggesting that intratumoral distribution of the targeted nanoparticles varies in different areas (Figure 6C). Interestingly, areas with the greatest decline in signal were largely in the periphery regions of the tumor mass which are rich in blood vessels, relative to the necrotic areas at the center of the tumor.
Molecular imaging probes targeting specific cancer markers have been long sought-after in applying tumor imaging for disease-specific detection and personalized therapeutics. However, the development of receptor-targeted imaging and its in vivo applications are particularly challenging, with current obstacles including: 1) the identification of cell surface biomarkers that are expressed sufficiently in tumor cells or tumor environments for sensitive tumor imaging; 2) the production of stable and high affinity targeting ligands in large amounts for in vivo studies; and 3) the development of safe and biodegradable contrast agents producing strong imaging signal or contrast.
The uPAR-targeted IO nanoparticle probe reported here provides an example that addresses these challenges and demonstrates the feasibility of in vivo receptor-targeted tumor imaging. Results of our study showed that ATF-IO nanoparticles are capable of targeting uPAR-expressing tumor cells in vitro and in vivo, and enable receptor-targeted MR and optical imaging of the tumor in vivo. We believe that the following characteristics of ATF-IO nanoparticles enabled uPAR-targeted imaging in vivo. First, we used a tumor targeting ligand from a natural high affinity receptor binding domain of uPA. uPA is composed of three independently folded domain structures: growth factor domain (GFD), Kringle domain, and serine protease domain. uPA binds with high affinity to uPAR through the ATF of the GFD, with a Kd less than 1 nM (39). Studies have shown that ATF (residues 1–135 aa) of uPA is a potent antagonist of uPA/uPAR binding since ATF lacks the serine protease domain of uPA, which negatively regulates its function by cleaving uPAR into a non-binding receptor (40, 41). Second, efficient internalization of the ligand/receptor complex may increase the concentration of the IO nanoparticles in tumor cells, which enhances the effect of uPAR-targeted tumor imaging (35, 42). Third, nanoparticles provide favorable pharmacokinetics by prolonging blood circulation time, allowing sufficient amounts of nanoparticles to reach the tumor. Furthermore, our ability to produce the recombinant protein in a large scale is essential for preclinical and eventually clinical studies. It should be mentioned that this study was done using mouse ATF peptides and the 4T1 mouse tumor model to investigate the feasibility of targeting uPAR. Although it has been shown that the interaction of uPA with its receptor has a species specificity (43), we found that mouse ATF peptides bind efficiently to mouse tumor cells and also show cross reactivity with human uPAR-expressing tumor cells. A major advantage of using mouse ATF peptides to conduct studies in a mouse tumor model is that the targeting specificity and biodistribution in normal tissues of this imaging probe can be studied in greater detail. For future clinical application, it is desirable to use a human ATF peptide. We have produced recombinant human ATF peptides and have demonstrated target specificity in human breast cancer xenograft models in nude mice (unpublished results).
Extensive studies have shown that human breast cancer and tumor stromal cells have a much higher level of uPAR compared to normal breast tissues (44, 45). Differences between the level of uPAR present on the surface of normal and tumor cells suggest that an elevated uPAR in cancer may be sufficient for in vivo uPAR-targeted tumor imaging. Several studies have shown that the highest level of uPAR expression is detected in the invasive edge of the tumor regions (23, 28), which are usually enriched in blood vessels, making this area particularly accessible for uPAR-targeted IO nanoparticles. Additionally, the ability of nanoparticles to leak cross tumor endothelium but not normal vasculatures by passive targeting may prevent the interaction of targeted nanoparticles with some cell types in normal tissues that express a low level of uPAR.
In this study, we prepared high quality and uniformly sized IO nanoparticles with a thin amphiphilic copolymer coating (estimated at ~2 nm). Compared with conventional dextran or PEG-coated nanoparticles used previously, amphiphilic copolymer-coated IO nanoparticles form a relatively small particle complex (~18 nm, in this study), which is desirable for in vivo delivery of the imaging probe. Several previous studies used a commercially available superparamagnetic iron oxide (SPIO) nanoparticle, Feridex, which has a particle core size distribution of 20–30 nm or 80–150 nm in overall diameter with dextran coating (13). We believe that size uniformity is essential for potential quantification using MRI signal of the amount of probe in vivo.
Although other small molecule imaging agents may have better intratumoral distribution than nanoparticle-based imaging agents, these imaging agents are usually eliminated from the blood circulation in a relatively short time (less than 30 min), which makes it unlikely that sufficient levels of the targeted contrast agents can accumulate at the tumor site (2). It has been shown that polymer-coated IO nanoparticles have over 8 hours of plasma retention time (11). This longer circulation time could be an important factor enabling targeted IO nanoparticles to reach the tumor site and bind to tumor cells. ATF-IO nanoparticles were stable in vivo and in intracellular environments for over 48 hrs during our imaging experiments. We observed that the intratumoral NIR signal increased over time and reached its highest level around 48 hrs after the administration of IO nanoparticles, suggesting that the long blood retention time may facilitate nanoparticle tumor targeting.
It has been reported by several other groups that large proportions of magnetic IO nanoparticles are taken up by the reticuloendothelial system in the liver and spleen, and then are subsequently metabolism or utilized for iron storage (12, 13). Our data showed that ATF-IO nanoparticles have reduced liver and spleen uptake compared with non-targeted IO nanoparticles. This suggests that conjugation of ATF-peptides to the IO nanoparticles attenuates their non-specific capture and retention in the liver and spleen, which commonly occurs after systemic delivery.
In conclusion, we have developed a uPAR-targeted molecular imaging nanoprobe that has a uniform-sized IO nanocrystal core, a thin amphiphilic copolymer coating, and a high affinity receptor binding domain of uPA conjugated with an NIR dye. This receptor-targeted nanoprobe selectively binds to and is internalized by tumor cells and can specifically accumulate in primary and metastatic tumors, facilitating in vivo MR and optical imaging in a mouse mammary tumor model. Such uPAR-targeted imaging nanoparticles are promising probes for molecular MRI of breast cancer and several other cancer types, such as pancreatic, lung and brain, which express high levels of uPAR (46–48). Given that chemotherapy drugs can be incorporated into the targeted nanoparticles (49), it will be feasible to use uPAR-targeted IO nanoparticles for delivery of therapeutic agents into tumor cells and monitoring the response to therapy using MRI.
Breast cancer is the most common cancer in women with about 180,000 new cases and 40,000 deaths each year in the United States. Currently, breast magnetic resonance imaging (MRI) is used for early detection of breast cancer and is recommended as a routine screen for women at high risk of the disease. Although breast MRI using a non-targeted contrast agent has a relatively high sensitivity in detecting lesions, it lacks specificity in determining the pathological characteristics of the lesions. Recent studies show that patients receiving MRI have a higher rate of biopsy and mastectomy compared to those without this imaging procedure. MRI is a promising non-invasive imaging approach for preoperative staging of breast cancer and monitoring tumor response to therapy. Therefore, the development of biomarker-targeted MRI contrast probes that increase the specificity of breast cancer MRI should have a great impact on the detection and treatment of breast cancer.
This research project is supported by the Emory-Georgia Tech Nanotechnology Center for Personalized and Predictive Oncology of the NIH NCI Center of Cancer Nanotechnology Excellence (CCNE, U54 CA119338), NIH NCI R01 CA133722 (LY), Emory Molecular and Translational Imaging Center grant (NIH, NCI, P50CA128613), Seed Grant from EmTech Bio, Inc (HM), the Friends for An Early Breast Cancer Test Foundation, the Idea Award of the Breast Cancer Research Program of the Department of Defense (BC021952), and the Nancy Panoz Endowed Chair (LY)
We would like to thank Dr. Anthea Hammond for her critical editing of the manuscript and Dr. Mark W. Dewhirst for kindly providing us with luciferase gene stable 4T1 cell line. We also thank Drs. Adam Marcus and Katherine Schafer-Hales in the Cell Imaging Core of the Winship Cancer Institute for their assistance in confocal microscopy, and Dr. Hongwei Duan for his help with zeta potential and particle size analysis.