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. 2010 Jun;24(6):1689–1699. doi: 10.1096/fj.09-139865

RGD peptide functionalized and reconstituted high-density lipoprotein nanoparticles as a versatile and multimodal tumor targeting molecular imaging probe

Wei Chen *, Peter A Jarzyna *, Geralda A F van Tilborg ‖, Van Anh Nguyen †, David P Cormode *, Ahmed Klink *,¶, Arjan W Griffioen #, Gwendalyn J Randolph †, Edward A Fisher **, Willem J M Mulder *,†,1, Zahi A Fayad *,‡,§,1
PMCID: PMC2874482  PMID: 20075195

Abstract

High density lipoprotein (HDL), an endogenous nanoparticle, transports fat throughout the body and is capable of transferring cholesterol from atheroma in the vessel wall to the liver. In the present study, we utilized HDL as a multimodal nanoparticle platform for tumor targeting and imaging via nonspecific accumulation and specific binding to angiogenically activated blood vessels. We reconstituted HDL (rHDL) with amphiphilic gadolinium chelates and fluorescent dyes. To target angiogenic endothelial cells, rHDL was functionalized with αvβ3-integrin-specific RGD peptides (rHDL-RGD). Nonspecific RAD peptides were conjugated to rHDL nanoparticles as a control (rHDL-RAD). It was observed in vitro that all 3 nanoparticles were phagocytosed by macrophages, while αvβ3-integrin-specific rHDL-RGD nanoparticles were preferentially taken up by endothelial cells. The uptake of nanoparticles in mouse tumors was evaluated in vivo using near infrared (NIR) and MR imaging. All nanoparticles accumulated in tumors but with very different accumulation/binding kinetics as observed by NIR imaging. Moreover, confocal microscopy revealed rHDL-RGD to be associated with tumor endothelial cells, while rHDL and rHDL-RAD nanoparticles were mainly found in the interstitial space. This study demonstrates the ability to reroute HDL from its natural targets to tumor blood vessels and its potential for multimodal imaging of tumor-associated processes.—Chen, W., Jarzyna, P. A., van Tilborg, G. A. F., Nguyen, V. A., Cormode, D. P., Klink, A., Griffioen, A. W., Randolph, G. J., Fisher, E. A., Mulder, W. J. M., Fayad, Z. A. RGD peptide functionalized and reconstituted high-density lipoprotein nanoparticles as a versatile and multimodal tumor targeting molecular imaging probe.

Keywords: angiogenesis, optical imaging, magnetic resonance imaging

High-density lipoprotein (HDL) is an endogenous nanoparticle, which is involved in lipid transport and cholesterol efflux in the body (1,2,3). As an endogenous nanocarrier, HDL nanoparticles are compatible with the immune system and can escape removal by the reticuloendothelial system. They naturally consist of a hydrophobic core (triglycerides and cholesteryl esters) surrounded by a monolayer of phospholipids embedded with cholesterol and apolipoprotein components. Their size is in the 7- to 15-nm range, and they are stabilized by the primary protein component apolipoprotein A-I (apoAI). Reconstituted HDL (rHDL) nanoparticles can be either discoidal (without the core) or spherical (with the core). Different targeting moieties and diagnostic/therapeutic agents may be included into the lipid layer or hydrophobic core or conjugated to apoAI of HDL nanoparticles. All the aforementioned features of the HDL nanoparticle platform make it a promising biocompatible and versatile multifunctional platform for drug and contrast agent delivery (4).

The natural interaction of HDL with macrophages facilitates its use as a molecular imaging agent in atherosclerosis. The first reported HDL imaging agent was a radioactive iodinated compound that was shown to accumulate in the atherosclerotic aortas of apolipoprotein E knockout (apoE−/−) mice (5). We previously developed both discoidal and spherical rHDL nanoparticles that contained paramagnetic lipids for magnetic resonance (MR) imaging (6, 7). The application of HDL MR contrast agents caused enhanced MR intensity in the macrophage-rich areas of aortic plaques in apoE−/− mice. In a subsequent study, the incorporation of a lipopeptide into the lipid layer of discoidal rHDL was demonstrated to further enhance their uptake by intraplaque macrophages (8). HDL has also proven to be a useful vehicle for the delivery of anticancer drugs to tumors (4). However, these HDL targeting strategies have thus far only resulted in imaging of macrophages, a natural target of HDL nanoparticles.

Recently, Zheng et al. (9, 10) developed a sophisticated method to reroute low-density lipoprotein (LDL) nanoparticles from their natural LDL receptors to the folate receptor by conjugating folic acid to the protein components of LDL. The rerouted LDL nanoparticles were used for optical imaging of cancers overexpressing the folate receptor (9,10,11). With the use of the same strategy, HDL nanoparticles were also rerouted to the folate receptor for optical imaging of tumors (11). Although optical imaging has high sensitivity and temporal resolution, its poor spatial resolution, the limited penetration depth of light, and lack of anatomical definition restrict its application as a deep tissue molecular imaging technique. On the other hand, MR imaging exhibits a better spatial resolution on anatomy and depicts opaque soft tissues with excellent contrast. However, MR imaging suffers from a low inherent sensitivity. Therefore, a dual modality HDL contrast agent that can be used for both the above-mentioned complementary in vivo imaging techniques will enhance the accuracy of molecular imaging using this probe.

In the present study, we developed an HDL nanoplatform that allows molecular imaging of tumors with both optical and MR imaging. To this aim, we synthesized discoidal rHDL containing a high payload of amphiphilic gadolinium chelates and a near infrared (NIR) fluorescent dye. Importantly, we demonstrated the ability to modify these rHDL nanoparticles to target angiogenic endothelial cells in tumors by conjugating αvβ3-integrin-specific cyclic 5-mer RGD peptides. The specificity of the HDL probes was evaluated in vitro in human umbilical vein endothelial cells (HUVECs) and murine J774A.1 macrophage cells. A human xenograft nude mouse model was used for in vivo targeting and imaging experiments, while ex vivo confocal microscopy was applied to corroborate and clarify the in vivo findings.

MATERIALS AND METHODS

Materials

1,2-Dipalmitoyl-sn-glycero-3-phosphocholine (DPPC) and 1,2-dimyristoyl-sn-glycero-3-phosphoethanolamine-N-(lissamine rhodamine B sulfonyl) (rhodamine-PE) were obtained from Avanti Polar Lipids (Albaster, AL, USA). Gadolinium diethylenetriaminepentaacetate-di(stearylamide) (Gd-DTPA-DSA) was purchased from IQSynthesis (St. Louis, MO, USA). The NIR dye, 1,1′-dioctadecyl-3,3,3′,3′-tetramethylindotricarbocyanine iodide (DiR), was purchased from Invitrogen (Carlsbad, CA, USA). N-succinimidyl-3-(2-pyridyldithio)-propionate (SPDP) was purchased from Pierce, Thermo Fisher Scientific (Rockford, IL, USA). The cyclic 5-mer RGD {c[RGDf(S-acetylthioacetyl]K} and cyclic 5-mer RAD {c(RADf[-S-acetylthioacetyl]K} were synthesized at a purity of 95% by Peptides International (Louisville KY, USA). RAD serves as a control in which one amino acid is different from RGD. The HEPES buffer contained 2.38 g/L HEPES (C8H18N2O4S, >99.5%; Sigma-Aldrich) and 8.0 g/L NaCl, and the pH was adjusted to 6.7 by addition of a NaOH solution. All other chemicals were of analytical grade or the best grade available. The chemical structures of Gd-DTPA-DSA and DiR are shown in Fig. 1.

Figure 1.

Figure 1.

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A) Chemical structures of Gd-DTPA-DSA, DiR, and RGD. B) Schematic of RGD peptide conjugation to discoidal rHDL nanoparticles. First, a linker was attached to the amine groups of exposed lysine units on apoAI. The activated RGD peptide was then covalently linked to rHDL via the linker.

Preparation of rHDL nanoparticles

Gd-DTPA-DSA, an amphiphilic fluorescent dye (either DiR or rhodamine-PE), and DPPC were dissolved in chloroform/ethanol (3:1 mol/mol) and initially dried under airflow at room temperature and subsequently under vacuum overnight. The dried lipid film was hydrated with a sodium cholate buffer (0.15 M NaCl and 1 mM EDTA, pH 7.5) at 55°C. Next, an apoAI (purified from human HDL) solution in PBS was added to the lipid solution that had been cooled to 4°C on ice. The molar ratio for apoAI:DPPC:Gd-DTPA-DSA:fluorescent dye:sodium cholate was 1:123:25:2:200. After incubation in an ice-water bath for 2 h, the mixture was dialyzed against buffer (0.15 M NaCl and 1 mM EDTA, pH 7.5) to remove the sodium cholate. The product was then filtered through a 0.22-μm syringe filter (Fisherbrand; Fisher Scientific, Pittsburgh, PA, USA) and concentrated, and the buffer was exchanged to 1% sucrose with a Vivaspin 6 centrifugal filter device (membrane cutoff range: 10,000 kDa; Sartorius Corporation, Edgewood, NY, USA).

Preparation of cyclic 5-mer RGD peptide conjugated rHDL (rHDL-RGD) nanoparticles

The conjugation of RGD peptides to rHDL nanoparticles is schematically illustrated in Fig. 1. The rHDL nanoparticles were functionalized with a SPDP linker before conjugation of the RGD peptides. First, the rHDL nanoparticles were transferred into HEPES buffer and diluted to 1 mg/ml apoAI, into which 20 mM SPDP in dimethylformamide was added in a 1:20 apoAI:SPDP ratio. After incubation at room temperature for 2 h, the reaction solution was thoroughly washed with a HEPES buffer using a Vivaspin 6 centrifugal filter device (membrane cutoff range: 10,000 kDa). The cyclic 5-mer RGD peptide was deacetylated in 0.05 M HEPES/0.05 M hydroxylamine-HCl/0.03 mM ethylenediamine tetraacetic acid (pH 7.0) for 1 h at room temperature. Next, the activated peptides were added to the rHDL-SPDP in HEPES buffer. This preparation was stored at 4°C overnight and purified using a Vivaspin 6 centrifugal filter device (membrane cutoff range: 10,000 kDa) by exchanging the HEPES buffer to PBS.

Nonspecific control nanoparticles (rHDL-RAD) were prepared by the same procedure using cyclic 5-mer RAD peptides.

Size and ζ-potential determinations

The mean sizes of the different nanoparticles were determined by photon correlation spectroscopy performed using a Malvern HPPS light-scattering instrument (Malvern Instruments, Malvern, UK). The ζ potential of the different nanoparticles was measured with a ZetaPALS analyzer (Brookhaven Instruments, Holtsville, NY, USA). All samples were analyzed at 25°C by adding 10 μl of sample in 1 ml filtered double-deionized water.

Gadolinium content and relaxivity determinations

The Gd concentrations of the rHDL-based nanoparticles were measured by inductively coupled plasma mass spectrometry. The relaxivity measurements were done on a Minispec (Bruker Medical, Ettingen, Germany) operating at 60 MHz and 40°C.

In vitro targeting of rHDL-RGD, rHDL-RAD, and rHDL-RGD

Murine macrophage J774A.1 cells (American Type Culture Collection, Manassas, VA, USA) were cultured in cell culture flasks (BD Falcon; BD Biosciences, San Jose, CA, USA) at 37°C in a 5% CO2 atmosphere using DMEM (Cellgro; Mediatech, Manassas, VA, USA) with 2 mM glutamine, 100 U/ml penicillin, 0.1 mg/ml streptomycin, and 10% FBS (Mediatech, Manassas, VA, USA). HUVECs were cultured in EGM medium (Clonetics; Lonza, Walkersville, MD, USA) supplemented with EGM-2-MV SingleQuots (Clonetics), 100 U/ml penicillin, and 0.1 mg/ml streptomycin. HUVECs were cultured up to the 4th passage. J774A.1 cells and HUVECs were grown to 60–70% confluency in 6-well plates. Cells were incubated for 1 or 24 h at 37°C with either the rHDL, rHDL-RGD, or rHDL-RAD nanoparticles at a concentration of 0.05 mM Gd. For competitive inhibition experiments, a 10-fold excess of free cyclic 5-mer RGD peptides in PBS was added to the cell culture medium 5 min before incubation with rHDL-RGD nanoparticles. After incubation, cells were washed with PBS and lysated in 500 μl Promega passive lysis buffer (Promega, Madison, WI, USA) for 30 min at room temperature. The collected cell lysates were centrifugated at 4000 rpm for 5 min to remove cell debris. The fluorescence intensity of rhodamine-PE of the lysates was measured using a BioTek Synergy 2 microplate reader with Gen5 software using a 540 ± 20 nm excitation filter and a 590 ± 20 nm emission filter (BioTek Instruments, Inc., Winooski, VT, USA). The protein concentrations of lysates were measured by the Bradford assay (12; Bio-Rad Protein Assay; Bio-Rad, Hercules, CA, USA) according to the protocol.

For MR imaging of cell pellets, macrophages and HUVECs at 60–70% confluence in cell culture plates were incubated for 1 or 24 h at 37°C as described above. Subsequently, the cells were washed twice with PBS, harvested, collected in 15-ml falcon tubes, and washed twice in PBS by centrifugation; 4% paraformaldehyde solution was added to the cells to make a final volume of ∼200 μl. These cell solutions were collected in small PCR cups and left overnight to form loosely packed cell pellets. MR imaging was performed on preparations of ∼106 packed cells.

The cups containing the cell pellets were placed in a custom-made sample holder, capable of carrying 3 PCR cups. A T1-weighted multislice multiecho (MSME) sequence (TR/TE=800 ms/10.5 ms, FOV=2.0×2.0 cm, matrix size=128×128, number of average=8, scan time=20 min, and slice thickness=0.5 mm) was used to generate T1-weighted images of the pellets.

Tumor mouse model

The Institutional Animal Care and Use Committee of Mt. Sinai School of Medicine approved the animal protocols and procedures. Human EW7 Ewing’s sarcoma cells were maintained in RPMI 1640 medium supplemented with 10% FCS. The cells were grown in a 5% CO2 and water-saturated atmosphere at 37°C, and subculturing was performed 1×/wk by 1:10 dilution after trypsinization. To generate tumor xenografts, ∼2 × 106 EW7 cells were injected subcutaneously into the right flank of the Swiss nude mice. Between d 21 and 28, when tumors had grown to a diameter of 4–5 mm, mice were used for imaging.

In vivo imaging

For in vivo NIR imaging, mice were anesthetized with a 4% isoflurane/O2 gas mixture (400 ml/min initial dose) and maintained with a 1.5% isoflurane/O2 gas mixture (100 ml/min maintenance dose) delivered through a nose cone. The nanoparticles (1 μmol of DiR/kg; n=5 mice/group) were intravenously administrated using a catheter that was placed in the tail vein. NIR images were collected using a custom-made imaging system (13, 14) with a 760-nm excitation filter, 800-nm cutoff emission filters, and 800 ms/frame. The system was controlled with LabVIEW (National Instruments, Austin, TX, USA; ref. 15).

For in vivo MR imaging, the animals were anesthetized with a 4% isoflurane/O2 gas mixture (400 ml/min initial dose) and maintained with a 1.5% isoflurane/O2 gas mixture (100 ml/min maintenance dose) delivered through a nose cone. An infusion line with the contrast agent was brought into the tail vein. Animals were imaged on a 9.4-T MRI scanner (Bruker Instruments, Billerica, MA, USA). The animals were anesthetized with a 4% isoflurane/O2 gas mixture (400 ml/min initial dose) and maintained with a 1.5% isoflurane/O2 gas mixture (100 ml/min maintenance dose) delivered through a nose cone and positioned in a 30-mm birdcage coil. A respiratory sensor connected to a monitoring and gating system (SA Instruments, Stony Brook, NY, USA) was placed on the abdomen to monitor the depth and frequency of respiration. At the beginning of MR imaging, tumors were localized using a T2-weighted multislice multiecho sequence (TR/TE/TE=2000 ms/30 ms/60 ms, FOV=2.56×2.56 cm2, matrix size=256×256, 21 contiguous 1-mm-thick axial slices, number of averages=8, and total scan time=15 min). High resolution T1-weighted images were generated using a spin echo (SE) sequence (TR/TE=800 ms/9.5 ms, FOV=2.56×2.56 cm2, matrix size=256×256, 21 contiguous 1-mm-thick axial slices, number of averages=8, and total scan time=20 min). After precontrast imaging, the animals were administered the different HDL contrast agents (25 μmol of Gd/kg; n=4 mice/group) via the tail vein in the magnet. Postcontrast imaging parameters were cloned exactly from precontrast imaging and were collected on each animal at 1, 2, and 24 h after injection.

Imaging analysis

For in vivo NIR images, the signal intensity of the tumor was normalized to the signal intensity of the skin by the following equation: NER = (Itumor − Iskin)/Iskin × 100%, where NER is the normalized enhancement ratio, Itumor is the real-time signal intensity of tumor, and Iskin is the real-time signal intensity of skin.

For in vivo MR images, a whole-tumor-based analysis was used to calculate the enhanced fractions, which did not require the slices or pixels to be exactly matched between pre- and postcontrast agent administration (16). The signal intensity of each single pixel was compared with a threshold derived from precontrast signal intensity of the tumor tissue. The tumor tissue region was determined in the T2-weighted images. The threshold was defined by: threshold = ItumorPre + 5 × Inoise, where ItumorPre is the mean signal intensity of tumor tissue in T1-weighted precontrast image, and Inoise is the mean signal intensity of noise in T1-weighed precontrast images divided by 1.25. Pixels within the tumor tissue were considered to be significantly enhanced when their signal intensity was higher than the threshold. Therefore, the precontrast image can also contain some pixels that are considered as enhanced according to this definition. The enhanced fraction within the whole tumor for each individual mouse was determined by the number of enhanced pixels divided by the number of total pixels within the tumor. A Matlab R2007b program (MathWorks, Natick, MA, USA) was used to do the routing imaging analysis.

Immunohistochemistry

Alexa Fluor 488-labeled isolectin GS-IB4 (a general stain for vascular endothelium; Invitrogen) was administered intravenously 10 min before the mice were sacrificed (17,18,19). The mice were perfused with PBS to remove unbound isolectin and/or nanoparticles. Tumor tissues were then removed and embedded in TissueTek (Sakura, Torrance, CA, USA). Eight-micrometer frozen sections were fixed in 4% paraformaldehyde and washed twice in PBS. The tissue sections were permeabilized in PBS with 0.5% Triton X-100 at room temperature for 30 min and washed with PBS twice. The sections were blocked in 2% goat serum in PBS for 45 min at room temperature. After blocking, the sections were incubated in 2% goat serum in PBS with Alexa Fluor 647-conjugated anti-mouse CD68 antibody (1:10 dilution; Serotec, Raleigh, NC, USA) overnight at 4°C and washed by PBS twice. Finally, the sections were mounted with DAPI-containing Vecta Shield mounting medium (Vector, Burlingame, CA, USA) and sealed with coverslips, shielded from light, and kept at 4°C until laser-scanning confocal fluorescence microscopy imaging was performed within 48 h. Sections without any antibody were used as negative controls. Confocal imaging was performed using a Leica SP5DM microscope (Leica Microsystems, Wetzlar, Germany). The system is equipped with four lasers: a 405-nm blue diode, an argon (488 nm), a green HeNe (543 nm), and a red HeNe (633 nm).

To examine the vessel density and distribution, the frozen tumor sections were fixed with cold acetone, blocked using serum (Vectastain ABC-AP kit; Vector), and then stained using a rat anti-mouse CD31 primary antibody (1:50; BD Pharmingen, San Jose, CA, USA) for 30 min at room temperature. After being washed, the sections were incubated with anti-rat IgG biotinlated secondary antibody (Vectastain ABC-AP kit) for 30 min at room temperature, followed by avidin-alkaline phosphatase (Vectastain ABC-AP kit) and Vector red alkaline phosphatase substrate kit (Vector). Finally, the sections were countedstained by Vector hematoxylin QS (Vector), dried, and mounted in Vector mounting medium. Each section was imaged at ×2 for a whole-tumor view. Microvessel density was assessed by counting randomly chosen areas at ×4 for 3 sections in each group.

Statistics

Data are presented as means ± se. For differences between groups, 2-way ANOVA was used. Multiple comparisons with Bonferroni correction after ANOVA were used to distinguish the groups of significant difference. A value of P < 0.05 was considered statistically significant. The analysis was performed on Matlab R2007b.

RESULTS

Physical properties

rHDL-RGD nanoparticles with DiR dye had a mean diameter of 12.1 ± 2.1 nm in aqueous solution as determined by dynamic light scattering, which was similar to and not significantly different from the size of apoAI-containing rHDL (rHDL) nanoparticles (9.3±1.7 nm) or rHDL-RAD nanoparticles (9.5±0.9 nm). At 60 MHz, the longitudinal relaxivity (r1) values of the nanoparticles were 8.5 ± 0.7, 8.7 ± 0.2, and 8.4 ± 0.6 mM−1·s−1 for rHDL, rHDL-RGD, and rHDL-RAD nanoparticles, respectively. In comparison, the relaxivity of the Gd-DTPA (Magnevist) is ∼3 mM−1·s−1 at this field strength. These physical properties of rHDL were consistent with our previous reports (6, 7). The conjugation of peptide to the exposed lysine units on apoAI changed the ζ potential of the nanoparticle. The ζ potentials of rHDL, rHDL-RGD, and rHDL-RAD nanoparticles in PBS were −12.7 ± 2.5, −36.4 ± 2.9, and −56.7 ± 0.8 mV, respectively.

rHDL-RGD targeting proliferating endothelial cells in vitro

The association of rHDL, rHDL-RGD, and rHDL-RAD with HUVECs and macrophage J774A.1 cells was compared by confocal microscopy (Fig. 2). After a 1 h incubation with the nanoparticles, a significantly higher level of association of fluorescently labeled rHDL-RGD nanoparticles with proliferating HUVECs in culture was observed, as compared with HUVECs incubated with rHDL or untargeted control rHDL-RAD nanoparticles. The competition inhibition experiment with free cyclic 5-mer RGD peptide remarkably decreased the association of rHDL-RGD nanoparticles with HUVECs. These results demonstrated the specificity of rHDL-RGD nanoparticles targeting proliferating endothelial cells in vitro. On the other hand, we did not observe association of any of these nanoparticles with macrophages after 1 h incubation by microscopy. After incubation for 24 h with nanoparticles, although associations with HUVECs were seen in all experiments, rHDL-RGD showed a much higher level of association with HUVECs than rHDL, rHDL-RAD, and rHDL-RGD under competitive inhibition conditions. A high level of association with macrophages was observed for all nanoparticles after 24 h incubation.

Figure 2.

Figure 2.

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Confocal microscopy images of HUVECs and murine J774A.1 macrophage cells incubated with medium only (control) or rHDL, rHDL-RGD, and rHDL-RAD nanoparticles for either 1 or 24 h. The competition inhibition experiments were performed with a 10-fold excess of free RGD (RGD compt.). All nanoparticles were labeled with rhodamine-PE (red). Nuclei were stained with DAPI (blue).

The uptake of nanoparticles was further quantified by measuring the fluorescence intensity of cell lysates. The fluorescence intensities were normalized to the protein concentration to correct the variation of cell numbers in each experiment. After 1 h incubation, the normalized fluorescence intensity of HUVECs with rHDL-RGD was 558 ± 63, while the values for rHDL, rHDL-RAD, and free RGD competition inhibition were 203 ± 59, 126 ± 68, and 234 ± 40, respectively (Fig. 3A). These results indicated that the rHDL-RGD interacted with cultured HUVECs specifically. At this time point, the macrophages did not show significant uptake in any of the incubations (Fig. 3B). After incubation for 24 h, increased association of nanoparticles with HUVECs was observed (Fig. 3C). The rHDL-RGD nanoparticles still showed a significantly higher level of association with HUVECs (1167±42) than rHDL (498±15), rHDL-RAD (538±19), and after competition with free RGD (675±36). Macrophages showed marked uptake of all 3 nanoparticles in comparison with HUVECs in terms of normalized fluorescence intensity (Fig. 3D). This quantitative difference of nanoparticle uptake by the 2 cell types at 2 different time points suggested different kinetics of binding/accumulation for different cell types and different nanoparticles in vitro.

Figure 3.

Figure 3.

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A–D) Normalized fluorescence intensity of cell lysates from HUVECs (A, C) and murine J774A.1 macrophage cells (B, D) incubated with medium only (control) or rHDL, rHDL-RGD, and rHDL-RAD nanoparticles for either 1 h (A, B) or 24 h (C, D). Competitive inhibition experiments were performed with a 10-fold excess of free RGD peptide (RGD compt.). All nanoparticles were labeled with rhodamine-PE (red). Bar graph represents means ± se. E) Representative T1-weighted MR images of cell pellets of HUVECs and murine J774A.1 macrophage cells incubated with nanoparticles under the same incubation conditions as in A–D. Total cell numbers were ∼1 × 106 for each pellet.

HUVECs and macrophage cell pellets were imaged using high-field MR. In the T1-weighted image (Fig. 3E), the pellets of HUVECs incubated with rHDL-RGD nanoparticles caused greater increase in MR signal enhancement relative to cells incubated with medium only (control), rHDL, rHDL-RAD, or free RGD competitive inhibition at either the 1 or 24 h time point. These results are indicative of a more pronounced association of rHDL-RGD specifically with HUVECs, as compared with all the controls. The macrophage cells showed strong MR signal enhancement for all nanoparticles after incubation for 24 h but not after incubation for 1 h. These MR imaging results of cell pellets are consistent with the results gained from confocal microscopy of cells and quantitative fluorescence measurements of cell lysates.

All rHDL nanoparticles caused MR signal enhancement of the tumor

All 3 rHDL-based nanoparticles loaded with amphiphilic Gd chelates were able to enhance the MR imaging signal in the tumor region, as observed in T1-weighted images (Fig. 4). A whole-tumor-based analysis was used to calculate the enhancement of tumor tissues, as described in Materials and Methods, that does not require the slices to be matched at the different time points. As expected, the preinjection images already showed some “enhanced” pixels, as their intensities were calculated to be above the threshold (Fig. 4A). Following injection of all 3 rHDL-based nanoparticles, the MR signal intensities of tumors (n=4 mice/group) showed a trend of enhancement at 1 and 2 h postinjection and were significantly enhanced at 24 h postinjection in comparison with the preinjection situation (Fig. 4B). The MR images also showed that the spatial distribution of enhanced pixels was not homogeneous in tumor tissues. However, no significant difference of MR enhancement was observed between the 3 rHDL-based nanoparticles.

Figure 4.

Figure 4.

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A) Representative in vivo T1-weighted MR images of Swiss nude mice bearing subcutaneous human EW7 Ewing’s sarcoma tumors. Whole-tumor-based analysis was used to calculate enhancement of MR signal. Enhanced pixels within the tumors were color coded. Gray scale represents signal intensity; color scale represents the percentage of enhanced pixels above a threshold that is defined by the mean intensity of the whole tumor and noise of precontrast MR scanning. B) Enhanced fraction of tumor tissue preinjection and at 1, 2, and 24 h postinjection of contrast agent. Bars represent means ± se (n=4 mice/group).

Enriching rHDL with RGD peptides accelerates the in vivo tumor binding/accumulation

NIR imaging was used to evaluate the optical imaging efficacy as well as to investigate the binding/accumulation kinetics of nanoparticles in the tumors in vivo. For this, a human sarcoma xenograft model (EW7 Ewing’s sarcoma, n=5 mice/group) was used. Figure 5A shows typical NIR images of mice bearing tumors before and up to 24 h after intravenous injection of nanoparticles. No autofluorescence was detected from the nude mice before injection as shown in preinjection images. The fluorescence signal was detected from the body of a mouse immediately after injection of nanoparticles, suggesting the rapid distribution of nanoparticles in the mouse body via the circulation. The rHDL-RGD nanoparticles bound/accumulated rapidly in the tumor region, while rHDL and rHDL-RAD nanoparticles showed slow binding/accumulation kinetics. In Fig. 5A, fluorescence from the tumor region was clearly greater than that in the surrounding skin at 30 min postinjection of rHDL-RGD nanoparticles, while the tumor regions were not observable for rHDL or rHDL-RAD nanoparticles at this time point. At 1 h postinjection, the fluorescence from the tumor region of mice injected with rHDL-RGD nanoparticles increased further and was remarkably higher than those injected with rHDL or rHDL-RAD nanoparticles. At 24 h postinjection, the tumor regions from all mice injected with any of the nanoparticles had a pronounced and similar NIR signal. Real-time quantitative analysis of NERs revealed significantly different binding/accumulation kinetics of the rHDL-RGD nanoparticles in comparison with rHDL and rHDL-RAD nanoparticles (Fig. 5B). rHDL-RGD nanoparticles caused a rapid increase of NER in tumors within 2 h, whereas injection of rHDL or rHDL-RAD nanoparticles resulted in a slow increase of NER in tumors during this time period. At 6 h postinjection and thereafter, the NERs of tumors reached similar levels for all 3 nanoparticles.

Figure 5.

Figure 5.

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A) Representative NIR images of Swiss nude mice bearing subcutaneous human EW7 Ewing’s sarcoma tumors intravenously injected with rHDL, rHDL-RGD, and rHDL-RAD nanoparticular contrast agents. Preinjection images and images at 5 and 30 min and 1, 2, and 24 h postinjection of contrast agent are shown. Arrows indicate tumor position. B) Normalized enhancement ratio of tumor region calculated from real-time signal intensities of the tumor region and of skin on the right leg. Data points on each line represent means ± se (n=5 mice/group).

rHDL-RGD nanoparticles colocalized with angiogenic endothelial cells

Ex vivo fluorescence confocal microscopy was used to investigate the binding/accumulation location of nanoparticles inside tumor tissues at 1 h (Fig. 6) and 24 h (Fig. 7) postinjection. At both time points (1 and 24 h postinjection), rHDL-RGD nanoparticles colocalized with endothelial cells within tumor tissues. At 24 h, rHDL-RGD nanoparticles were also observed to accumulate in the interstitial space of the tumors. However, rHDL and rHDL-RAD nanoparticles were only observed in the interstitial space of the tumor but not colocalized with endothelial cells at either 1 or 24 h postinjection. Colocalization of nanoparticles in the interstitial space with macrophages was not observed, although a fraction was occasionally found to be associated with CD68 stained macrophages. The presence of microvessels in tumor sections was confirmed by CD31-positive staining (Fig. 8A–C), which revealed an inhomogeneous distribution with higher density near the tumor periphery. The microvessel densities of tumors were similar for all 3 groups (Fig. 8D).

Figure 6.

Figure 6.

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Confocal microscopy images of tumor sections at 1 h postinjection of rHDL, rHDL-RGD, and rHDL-RAD nanoparticles labeled with rhodamine-PE (rhodamine, red). Angiogenic endothelial cells (Isolectin-A488, green), macrophages (CD68-A647, magenta), and nuclei (DAPI, blue) are shown.

Figure 7.

Figure 7.

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Confocal microscopy images of tumor sections at 24 h postinjection of rHDL, rHDL-RGD, and rHDL-RAD nanoparticles labeled with rhodamine-PE (rhodamine, red). Angiogenic endothelial cells (Isolectin-A488, green), macrophages (CD68-A647, magenta), and nuclei (DAPI, blue) are shown.

Figure 8.

Figure 8.

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A–C) Microscopy images of tumor sections stained with CD31 and counterstained with hematoxylin from mice injected with rHDL (A), rHDL-RGD (B), and rHDL-RAD (C) nanoparticles. View: ×2. D) Microvessel density quantification in n = 3 sections/group. No significant differences were observed between groups. Data points on each line represent means ± se.

DISCUSSION

In the present study, we demonstrated that rHDL-based nanoparticles can be used as multimodality contrast agents for tumor imaging with in vivo NIR, MR and ex vivo confocal imaging capabilities. The rHDL nanoparticles can be rerouted to target angiogenic endothelial cells by conjugation with a cyclic 5-mer RGD peptide specific for the αvβ3 integrin.

We provided evidence that rHDL-RGD uptake by proliferating HUVECs was specific in vitro (Figs. 2 and 3A) resulting in an enhanced MR signal of the rHDL-RGD treated cell pellets (Fig. 3B). The competitive inhibition experiments using free cyclic RGD peptide remarkably decreased the uptake of rHDL-RGD nanoparticles by HUVECs. This indicates that the interaction between rHDL-RGD and HUVECs is through the RGD peptide. The RGD peptide is developed to specifically target the αvβ3 integrin (20), which is expressed by proliferating HUVECs. In a number of studies, we have reported the use of this peptide to direct a variety of synthetic nanoparticles to angiogenically activated tumor blood vessels (16, 21,22,23,24). In the current study, we have functionalized an endogenous nanoparticle with this peptide that successfully directed the modified rHDL nanoparticles to HUVECs through the αvβ3-integrin. Unmodified HDL nanoparticles are found to interact with macrophages specifically via several receptors (18, 25, 26). This endogenous mechanism was utilized for rHDL nanoparticles to naturally target macrophages in the atherosclerotic plaque for molecular imaging purposes (6, 7, 27, 28). Differences in uptake efficiency by HUVECs and macrophages were found between the different rHDL-based nanoparticles. The uptake/binding of rHDL-RGD to HUVECs was already observable by confocal microscopy images after incubation for 1 h, while we did not observe much uptake of the control nanoparticles by macrophages at this time point (Fig. 2). The rapid binding of rHDL-RGD nanoparticles was also observed in vivo using NIR imaging (Fig. 5). The association of rHDL-RGD nanoparticles with macrophages in vitro can be due to several factors, including the expression of αvβ3 integrin on the macrophage surface, the nonspecific endocytosis/phagocytosis by macrophages, and/or the binding of rHDL-RGD nanoparticles to other integrins (e.g., αvβ5 integrin; ref. 20).

The combination of NIR and MR imaging exploits the complimentary features of both techniques and thus provides high sensitivity, an excellent temporal resolution, and a good anatomical definition with a high spatial resolution. The rHDL-RGD particles were found to accumulate/bind in the tumor tissue rapidly (within 2 h after injection) and gave good tumor to skin contrast under sensitive NIR imaging (Fig. 5) but only a marginal increase of the MR signal (Fig. 4). This rapid enhancement of tumor to skin contrast for rHDL-RGD nanoparticles was due to their rerouting to αvβ3 integrin because no significant differences were observed for tumor microvessel densities between the 3 groups. However, MR images provided anatomical information on the enhanced regions of tumors (Fig. 4), which showed inhomogeneous distribution of enhanced pixels in tumors. The fluorescent NIR signal is more directly related with the accumulation/binding concentration of contrast agents, subjected to the minor attenuation of tumor size (5–6 mm) and skin thickness (<0.5 mm). Thus, the quantified real-time NIR signal can provide valuable information about the kinetics of nanoparticle accumulation/binding (Fig. 5B). Conversely, the MR signal is less reliable as a quantitative measure and is influenced by many other factors, including compartmentalization within the cell (29,30,31), exchange of water across the vesicular membranes (30,31,32), size of the cytoplasmic vesicles (32), interaction with endogenous cellular components (31, 33), binding status to the extracellular matrix (29), and availability of exchanging water to Gd contrast agents (31, 33).

To maintain the pathological growth of tumors, angiogenesis is required to supply nutrients for tumor tissue via newly formed blood vessels (34, 35). Therefore, antiangiogenesis is an interesting strategy for therapeutic intervention (34). To evaluate the progress of angiogenesis and monitor the antiangiogenesis therapies, methods to image this process in living animals and humans are needed. rHDL-RGD nanoparticles specifically targeted angiogenic endothelial cells of tumors in vivo. Fluorescence microscopy revealed that rHDL-RGD nanoparticles mainly colocalized with tumor blood vessels at both 1 and 24 h postinjection (Figs. 6 and 7). It has been shown that these angiogenic blood vessels highly express the αvβ3 integrin, to which the cyclic 5-mer RGD peptide binds, a strategy that we have previously exploited in a number of molecular imaging studies (16, 21,22,23,24). Therefore, these results suggest a specific association between rHDL-RGD nanoparticles and the αvβ3-integrin expressing endothelial cells. The accumulation patterns in the tumors of rHDL and/or rHDL-RAD nanoparticles were remarkably different from rHDL-RGD nanoparticles. The rHDL and rHDL-RAD nanoparticles were found in the interstitial space within the tumor, which is due to the nonspecific extravasation of the nanoparticles. The extravasation of these 2 nanoparticles is most likely the reason for their accumulation in tumor and for the slow increase of NIR contrast and MR enhancement fraction of tumor at later time points. It has to be emphasized that the extravasated nanoparticles were not colocalized with CD68-stained macrophages in vivo (Figs. 6 and 7), although macrophages associated with all the 3 nanoparticles in vitro. This is most likely due to the diffusion/permeation limit within tumor tissue, the availability of nanoparticles to macrophages, and/or the small population of macrophages near the abnormal neovessels.

Lipoprotein nanoparticles, especially LDL, have been used for delivery of drugs to tumors (4) and have also been rerouted for tumor imaging (9,10,11, 36, 37). In comparison with LDL nanoparticles, the HDL-based nanoplatforms have several advantages. The smaller size of HDL nanoparticles makes them more permeable to vascular membranes (9). HDL can be easily reconstituted from natural and/or synthetic lipids/proteins/peptides (6,7,8,9, 27, 28). HDL has therapeutic potential for atherosclerosis (3, 38), while LDL is atherogenic (39, 40). Corbin et al. (11) rerouted spherical reconstituted HDL nanoparticles (containing cholesteryl ester) to the folate receptor for in vivo optical NIR imaging of tumors overexpressing this receptor. In our study, we introduced additional in vivo MR modality to discodial HDL nanoparticles and rerouted them to angiogenic endothelial cells. The small dense discodial HDL nanoparticles do not contain cholesterol/cholesteryl esters. They therefore have a great capacity for acquiring unesterified cholesterol from macrophages, similar to the behavior of endogenous pre-β-HDL particles (41, 42). The multimodality character of our platform enables complementary functions for optical and MR imaging with high sensitivity and high anatomical resolution.

In summary, this study demonstrates that rHDL-based nanoparticles are a versatile platform to image tumors in vivo with multimodal capability to combine complementary imaging techniques. The nonspecific accumulation of unmodified rHDL nanoparticles can give rise to optical and MR contrast within tumors at a late time point. Molecular imaging of angiogenesis can be achieved through rerouting the nanoparticles to endothelial cells by conjugation of a cyclic 5-mer RGD peptide specific for the αvβ3 integrin. Different accumulation/binding kinetics for rHDL and rHDL-RGD nanoparticles was observed by optical imaging.

Acknowledgments

Partial support was provided by U.S. National Institutes of Health (NIH)/National Heart, Lung, and Blood Institute grants R01 H-71021 and R01 H-78667 and NIH/National Institute of Biomedical Imaging and Bioengineering grant EB-009638 (to Z.A.F.). We thank CSL Behring (Bern, Switzerland) for the kind gift of apoA-I. Confocal microscopy was performed at the Mount Sinai School of Medicine Microscopy Shared Resource Facility and supported by NIH–National Cancer Institute grant 5R24 CA-095823-04, National Science Foundation grant DBI-9724504, and NIH grant 1 S10 RR-09145-01.

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