RGD-targeted paramagnetic liposomes for early detection of tumor: In vitro and in vivo studies

European Journal of Radiology 80 (2011) 598–606

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European Journal of Radiology journal homepage: www.elsevier.com/locate/ejrad

RGD-targeted paramagnetic liposomes for early detection of tumor: In vitro and in vivo studies Wei Li, Bo Su, Shuyan Meng, Lixia Ju, Linghua Yan, Yongmei Ding, Yin Song, Wei Zhou, Heyan Li, Liang Tang, Yinmin Zhao, Caicun Zhou ∗ Research Institute of Oncology, Tongji University Medical School, 507 Zhenmin Road, Shanghai 200433, PR China

a r t i c l e

i n f o

Article history: Received 3 September 2010 Accepted 3 January 2011 Keywords: Magnetic resonance imaging Liposomes Integrins Lipopeptide Gadolinium DTPA Nanoparticles

a b s t r a c t Magnetic resonance molecular imaging has emerged as a potential approach for tumor diagnosis in the last few decades. This approach consists of the delivery of MR contrast agents to the tumor by specific targeted carriers. For this purpose, a lipopeptide was constructed by using a cyclic RGD peptide headgroup coupled to palmitic acid anchors via a KGG tripeptide spacer. Targeted paramagnetic liposomes were then prepared by the incorporation of RGD-coupled-lipopeptides into lipid bilayers for specific bounding to tumor. In vitro, study demonstrated that RGD-targeted liposomes exhibited a better binding affinity to targeted cells than non-targeted liposomes. MR imaging of mice bearing A549 tumors with the RGDtargeted paramagnetic liposomes also resulted in a greater signal enhancement of tumor compared to non-targeted liposomes and pure contrast agents groups. In addition, biodistribution study also showed specific tumor targeting of RGD-targeted paramagnetic liposomes in vivo. Therefore, RGD-targeted paramagnetic liposomes prepared in the present study may be a more promising method for early tumor diagnosis. © 2011 Elsevier Ireland Ltd. All rights reserved.

1. Introduction Non-invasive imaging is essential for diagnosing and monitoring disease particularly tumors. Besides symptoms, clinical diagnosis is usually made with conventional X-ray, CT (computed tomography), and MRI (magnetic resonance imaging). For early detection of tumor, these imaging modalities play a key role in the diagnosis as the symptoms might be absent or masked. However, conventional imaging techniques have limited sensitivity and specificity. Thus, there is an urgent need in the field of tumor imaging for special techniques which will be able to detect early stages of tumor, even those less than 1 cm in diameter. By using imaging physiology and pathological processes at the molecular level in vivo, the emerging field of molecular imaging enables diagnosis of early tumor with increasing sensitivity in CT, MRI [1–3], scintigraphy [4] and optimal imaging [5]. Furthermore, advances in imaging technologies may lead to novel insights into the mechanisms of tumor growth and progress, as well as to the study of some important molecular changes such as the high expression of cell surface receptors. One of the critical factors during tumor growth and development is angiogenesis without which tumors cannot grow to a diameter greater than a few millimeters [6,7]. During the cascade

∗ Corresponding author. Tel.: +86 21 65115006. E-mail address: [email protected] (C. Zhou). 0720-048X/$ – see front matter © 2011 Elsevier Ireland Ltd. All rights reserved. doi:10.1016/j.ejrad.2011.01.051

of angiogenesis, there is upregulation of several specific cell surface receptors, such as ␣␯, ␤1 integrins [8] and VEGFR [9] (vascular endothelial growth factor receptor) which mediate vessel growth on activated tumor endothelial cells. Most previous studies investigating molecular imaging markers of tumor angiogenic blood vessels have paid particular attention to the ␣␯␤3-integrin [10,11], a critical molecule for neovessel formation. This integrin which is highly expressed on the surface of activated endothelial cells and tumor cells but not on the quiescent endothelial cells has been linked with cell invasion and migration. Like other integrins, ␣␯␤3-integrin also binds to extracellular matrix via the three amino acid sequence of RGD (arginine-glycineaspartic acid) [12]. To date, a variety of RGD-containing-peptides have been widely used for molecular imaging [13]. In the majority of cases, these RGD-based imaging agents [14] were developed by direct conjugation of the homing peptide to drug or tracers. Several studies have used ␣␯␤3-specific RGD conjugated with a radiolabel like 125 I [15], 18 F [16,17] to image ␣␯␤3-integrin expression by nuclear methods such as PET (positron emission tomography) and scintigraphy imaging. It has been proved that this approach of nuclear imaging with radiolabelled RGD peptides is very effective and sensitive. Nevertheless, the limited spatial resolution of these imaging methods prevents exact location of the tumor especially in early stages. Compared with the nuclear methods above, MRI has a better spatial resolution and can obtain a precise anatomical localization. Furthermore, safety and absence of radioactivity

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are also among the advantages of MRI in early tumor detection and diagnosis. However, MRI suffers from a relatively low inherent sensitivity which leads to weak contrast resolution in differentiating tumor from the neighboring normal tissue. To overcome the limitations, effective measures could be taken including the use of nanoparticles with a high payload of MR contrast agents [18] and the exploitation of targeted receptor-ligand mechanism for specifically delivering these nanoparticles to tumor region [2,19]. In this report, novel RGD-targeted paramagnetic liposomes were constructed. Fluorescence labelled liposomes was prepared to test the specificity of the targeted liposomes on proliferating HUVEC and A549 cells in vitro. In vivo experiments were performed in a subcutaneously implanted A549 human lung adenocarcinoma xenograft model to define the distribution and tumor uptake of the Gd. We also did MRI scans to evaluate the ability of the targeted paramagnetic liposomes as a potent MRI contrast agent for tumor diagnosis. 2. Materials and methods 2.1. Chemicals Gd-DTPA (gadolinium diethylenetriamine penta-acetic acid) was purchased from Bayer Schering Pharma AG (Magnevist, Berlin, German). Egg phosphatidylcholine (egg PC, C42 H82 NO8 P) and cholesterol (C27 H46 O) were from Lipoid GmbH (Ludwigshafen, Germany), methoxy poly (ethylene glycol)-distearoylphosphatidylethanolamine 2000 (mPEG2000DSPE) was from Avanti Polar Lipids (Alabaster, AL, USA). Lissamine TM rhodamine B 1, 2-dihexadecanoyl-sn-glycero3-phosphoethanolamine, triethylammonium salt (Rho-DHPE) was from Invitrogen corporation (Carlsbad, CA, USA). All reagents and solvents of analytical or HPLC grade were obtained from commercial sources. All peptides, conjugates were synthesized by Yishengyuan Corporation (Shanghai, China). 2.2. Methods 2.2.1. Synthesis of lipopeptides Lipid-conjugated peptide was constructed by a RGD peptide headgroup, a KGG spacer and 2 palmitic acid (pal) anchors. The RGD-coupled-lipopeptides were incorporated into liposomes by the pal anchors which were intercalated within the liposome bilayer. All these lipopeptides were synthesized using Fmoc solidphase synthesis chemistry by Yishengyuan Corporation (Shanghai, China). The purity of the product (>90%) was confirmed by high performance liquid chromatography (HPLC). 2.2.2. Preparation of liposomes Liposomes were prepared by thin film hydration methods. Egg PC, CHOL, mPEG2000-DSPE were mixed at a molar ratio of 1.85/1/0.15. For peptide-coupled targeted liposomes, RGDcoupled-lipopeptides were incorporated at about 3 ␮g/␮mol ratio respectively. 0.1 mol% Rho-DHPE was added for fluorescent study. The mixture was dissolved in chloroform and evaporated to dryness in a round bottom flask on a rotary evaporation at 37 ◦ C, and then the lipid film was further dried by vacuum drying over night. Lipopeptides were dissolved in DMSO firstly and then diluted in chloroform to make the DMSO final concentration under 1‰. For paramagnetic liposomes, the lipid film was subsequently hydrated with 10 mM Gd-DTPA solution with PBS (phosphatebuffered saline) solution, followed by 6 freeze–thaw cycles. The liposome suspension was extruded 10 times sequentially through 400, 200 and 100 nm-polycarbonate membranes (Avanti Polar Lipids Inc., Alabaster, USA). For fluorescent study, non-paramagnetic liposomes were prepared by PBS hydration.

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Unencapsulated lipopeptides and Gd-DTPA were removed by centrifugating three times at 100,000 × g through ultrafilters (10kD NWMC, Ultracel YM-10, Millipore Corp., MA, USA) in an Avanti J-30I Centrifuge (Beckman Coulter, Krefeld, Germany) for 60 min. Final Gd concentration of paramagnetic liposomes was ∼10 mM. The encapsulation efficiency (%) of Gd-DTPA in liposomes was evaluated by ultrafiltration method. 2.2.3. Liposome characterization The morphology of prepared paramagnetic liposomes was examined by using a JEM 1200 EX electron microscope (JEOL, Tokyo, Japan). Mean particle size and size distribution of liposomes were determined by dynamic light scattering on a MASTERSIZER 2000 Submicron Particle Sizer (Malvern, UK). To determine the Gd concentration, paramagnetic liposomes were lysed by 10% (v/v) Triton solution. The amount of Gd was measured by inductively coupled plasma-atomic emission spectrometry (ICP-AES, Varian Inc., Lexington, MA, USA). For MRI analysis, 2 ml of controls of PBS, prepared paramagnetic liposomes and Gd-DTPA with a final concentration of 10 mM were placed in Eppendorf tubes. The T1 relaxation values of solutions were measured on a 3T magnetic resonance scanner (3T Trio, Siemens, Erlangen, Germany) at ambient temperature. T1 values were obtained using an inversion recovery spin–echo sequence with 11 different inversion times, ranging from 23 ms to 9000 ms, the number of signal averages was 2, FOV was 220 mm × 220 mm collected into a matrix of 320 × 320. The T1 is reported as mean ± SD of the pixels in a circular region of interest (ROI) defined by the contour of the tubes. 2.2.4. In vitro release of Gd-DTPA from liposomes Gd-DTPA release patterns of paramagnetic liposomes and pure Gd-DTPA were evaluated using the dialysis method. 1 ml of different contrast agents was placed in a dialysis bag (MWCO 14000, Gene Bio-Application Ltd., Israel). Bags were immersed in 100 ml of release medium (PBS containing 45 g/L bovine serum albumin to mimic in vivo condition). While stirring the release medium using a magnetic stirrer at 300 rpm, samples (40 ␮l) were taken at predetermined time points from the release medium over a 24 h period and refilled with the same amount of the fresh medium. The amount of Gd in each sample was measured by ICP-AES. 2.2.5. Cell culture The A549 cells, a human lung carcinoma cell line and human umbilical vein endothelial cells (HUVEC) were gifts from Central Laboratory of Shanghai Pulmonary Hospital (Shanghai, China). Both cell lines were cultured in Dulbecco’s Modified Eagle Media (DMEM) supplemented with 10% fetal bovine serum (FBS) and 1% antibiotics (100 U/ml penicillin G and 100 U/ml streptomycin). Cells were maintained in a 37 ◦ C humidified incubator containing 5% CO2 . For all experiments, cells were seeded in 6 well cell culture plates until wells reached ∼70–80% confluence, ∼106 cells of passage 2–3 were used. 2.2.6. In vitro targeting of liposomes For fluorescent experiments, targeted or non-targeted fluorescent liposomes were first diluted by culture medium without serum to a final concentration of 1 ␮mol lipid/ml and then added to culture plate. Each group was quadruplicated. For binding competition experiments 10-fold of free RGD peptides were added into the medium 2 h before adding of targeted liposomes. Cells were incubated for 4 h at 37 ◦ C and subsequently washed twice with PBS to remove unbound fluorescent liposomes. For fluorescent microscopy, cells were fixed with 4% formaldehyde and immediately visualized using inverted microscope (Olympus CKX41, Olympus Corporation, Tokyo, Japan).

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Fig. 1. Structure of RGD-coupled lipopeptide and targeted liposome. (A) RGD-coupled lipopeptide was constructed by a RGD peptide headgroup, a KGG spacer and two palmitic acid anchors. (B) Lipopeptides were incorporated into lipid bilayers of PEG-coating liposomes. Gd-DTPA was entrapped in the aqueous core of liposomes.

For fluorospectrophotometer test, cells were washed, harvested by a trypsinization procedure and collected in tubes after incubation. Then 300 ␮l cell lysis buffer was added to lyse the harvested cells and the cell lysate was mixed with 200 ␮l MeOH and 1200 ␮l CHCl3 by vortex mixing. The mixture was centrifuged at 4000 × g for 10 min. Afterwards, the organic phase was collected and the fluorescence intensity was measured by a fluorospectrophotometer (970 RT, Shanghai Lengguang Technology Co., Ltd.). Paramagnetic liposomes at a final Gd concentration of ∼1 mM were also incubated with cells followed by nitric acid digestion at 65 ◦ C overnight. The amount of Gd in digestion samples was determined by ICP-AES. 2.2.7. Mouse tumor model To prepare the xenograft mouse model, 4-week-old female BALB/C nude mice were obtained from the Shanghai SLAC Laboratory Animal Co. Ltd. A549 cells were collected by trypsin digestion at exponential growth phase and 1 × 107 cells in 0.2 ml of serum-free culture medium were injected into the upper right flank region of each mouse at day 0. When the tumors grew to a size of 50–100 mm3 (around day 10–14), mice were used for experiments. These tumor bearing mice were randomized into three groups: (1) pure Gd-DTPA (Gd-DTPA), (2) non-targeted paramagnetic liposomes (NT-Gd-LP), (3) RGD-targeted paramagnetic liposomes (T-Gd-LP) for MRI scan (n = 4–5) or biodistribution study (n = 4/group/time point). After anesthesia by intraperitoneal injection of 10% urethane (m/v), mice were injected with paramagnetic liposomes or pure GdDTPA (0.1 ␮mol Gd/g body weight, ∼200 ␮l total) via the tail vein. At predetermined time points mice had an MRI scan or biodistribution study. The animal study protocol was reviewed and approved by Shanghai Commission on Care and Management of Experimental Animals and the animal study was in accordance with the National Institute of Health Guide for the Care and Use of Laboratory Animals. 2.2.8. MRI imaging and image analysis For in vivo study, MR imaging was performed on a GE Signa Excite HD 1.5 Tesla horizontal bore scanner (GE Healthcare, Wisconsin, USA). Tumor-bearing mice were anesthetized and placed in a home-built cradle. To collect baseline data, the tumor was firstly localized by a T2-weighted scan, thereafter all mice were scanned using T1-weighted spin-echo sequence (TR = 440 ms, TE = 17 ms, FOV = 10.0 cm × 10.0 cm, slice thickness = 2 mm, matrix = 512 × 256). Subsequently these mice

were injected with paramagnetic contrast agents. Mice were rescanned 30 min, 1 h, 2 h, 4 h, and 6 h after the injection using the same parameters as previously described. The tumor tissue and right hind limb muscle tissue were segmented by manually drawing a ROI (region-of-interest) of 3 mm × 3 mm around the area in the T1-weighted images. The signal intensities in the ROIs were measured before and after the injection of the contrast agent and the SER (signal enhancement rate/%) was determined according to the following formula [20]:

 SER % =

(SITumor /SIMuscle )Post − (SITumor /SIMuscle )Pre (SITumor /SIMuscle )Pre

 × 100%

The mean SER value was averaged for the evaluation both in terms of percentages. 2.2.9. Biodistribution and tumor uptake study Appropriate dose of paramagnetic liposomes or pure Gd-DTPA were administered to animals via tail vein. After injection, these mice were sacrificed at predetermined time point. Blood, tumor, muscle, liver, spleen, heart, lung and kidney samples were obtained, weighted, and then digested with nitric acid at 65 ◦ C overnight. The amount of Gd in these organs was determined by ICP-AES. 2.2.10. Statistical analysis Statistical analysis was performed using one-way analysis of variance at the p < 0.05 level. All data were expressed in the form of the mean ± standard deviation. 3. Results 3.1. Synthesis and characterization As shown in Fig. 1, RGD-coupled lipopeptide was composed of a RGD peptide headgroup, a KGG tripeptide spacer and 2 palmitic anchors. These lipopeptides were incorporated into the lipid bilayers to construct targeted liposomes with Gd-DTPA entrapped in the aqueous interior phase as paramagnetic contrast agents. Transmission electron microscopy of all paramagnetic liposomes was homogeneous in size and round or discrete in shape. The particle diameter distribution of targeted liposomes was shown in Fig. 2. The average particle diameter was 99.1 ± 9.2 nm. There were no significant differences in particle size, size distribution and Gd concentration between targeted and non-targeted paramagnetic

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Fig. 2. (A) Transmission electron microscopy image of the targeted liposomes. (B) Size distribution of liposome particles dispersed in PBS solution.

liposomes (Table 1). The entrapment efficiency of different paramagnetic liposomes was around 52.0 ± 2.4%. MRI studies of the paramagnetic liposomes were also performed. Collected MR images and T1 values showed both non-targeted and RGD-targeted paramagnetic liposomes to have comparable T1 relaxation with pure Gd-DTPA which was widely used for clinical use. 3.2. In vitro release of Gd Shown in Fig. 3 was the data of release experiments. Because of high water solubility, release of Gd-DTPA from pure Gd-DTPA solution (Magnevist) was rapid with nearly 100% Gd-DTPA released within 1 h. When entrapped into normal liposomes, the leakage was slowed down. Approximately 28% were released in the first 4 h, furthermore 80% within the next 24 h. The release of entrapped Gd-DTPA from PEGylated liposomes was the slowest. In the first 4 h only about 18% was released and up to ∼25% in 24 h. This demonstrated that the paramagnetic PEGylated liposomes could achieve more preferable release effect. No significant difference in the release pattern was observed between targeted and non-targeted liposomes, suggesting that the incorporation of lipopeptides into the phospholipid-bilayers did not affect the stability of liposomes.

liposomes were incubated with HUVEC or A549 cells. After incubation, cells were washed and studied by fluorescent microscopy. The images in Fig. 4(A and B) showed that RGD-targeted liposomes efficiently bind to the cell surface whereas no significant binding of non-targeted liposomes was observed. In addition, to investigate whether the interaction between targeted liposomes and cells was mediated through RGD, free RGD peptides were added prior to incubation of liposomes with cells to perform a competition experiment. It clearly showed that the binding between RGD-targeted liposomes and cells was inhibited after adding of free RGD peptides. The same results were also obtained from the fluorescence spectrophotometer test as shown in Fig. 4(C and D). In both cell lines, the relative fluorescence intensity of targeted liposomes groups was approximately 10 times of non-targeted liposomes groups. It indicated that RGD-targeted liposomes had significantly higher binding affinity when compared to non-targeted liposomes. Moreover, the binding could be competitively inhibited by free RGD peptides.

3.3. In vitro targeting of liposomes To demonstrate the specific binding of RGD-targeted liposomes to integrin ␣␯␤3 receptor, targeted and non-targeted fluorescent Table 1 Physical characteristics of paramagnetic liposomes. Phosphate-buffered saline (PBS), pure Gd-DTPA (Gd-DTPA), non-targeted paramagnetic liposomes (NT-Gd-LP) and RGD-targeted paramagnetic liposomes (T-Gd-LP).

Particle size Gd concentration (mmol/l) (nm)

T1 value (ms)

PBS

−

−

2565.32±36.67

Gd-DTPA

−

10.14±1.63

18.66±3.21

NT-Gd-LP

89.1±8.8

10.4±1.51

25.04±6.03

T-Gd-LP

99.1±9.2

10.56±1.32

28.54±4.92

Fig. 3. Release behavior of Gd-DTPA from different formulations. Pure Gd-DTPA (GdDTPA), non-targeted normal liposomes (Gd-LP), non-targeted PEGylated liposomes (PEG-Gd-LP) and RGD-targeted PEGylated liposomes (T-PEG-Gd-LP) were dialyzed using sink solution of PBS (pH 7.4) containing 45 mg/ml bovine serum albumin. The values are arithmetic means of three experiments.

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Fig. 4. (A and B) Fluorescence microscopy and (C and D) spectrophotometer test of HUVEC and A549 cells after incubation with rhodamine-labelled liposomes. (A) HUVEC and (B) A549 cells were incubated for 4 h with respectively non-targeted fluorescence liposomes (NT-F-LP), RGD-targeted fluorescence liposomes (T-F-LP) and competition of RGD-targeted liposome conjugates with free RGD peptides (+free RGD). Non-targeted fluorescence liposomes (dotted line), RGD-targeted fluorescence liposomes (solid line), with free RGD peptides (dashed line). The arrow and dotted-and-dashed line showed the excitation wavelength of rhodamine-DHPE.

In order to further evaluate in vitro binding activity, paramagnetic liposomes were incubated with cells. The cell concentration of Gd for those mixed with RGD-targeted liposomes was higher as compared to non-targeted liposomes and pure Gd-DTPA in both cell lines. There was statistically significant difference among various formulations. Data were shown in Table 2. 3.4. In vivo MRI imaging For MRI imaging, different formulations of paramagnetic contrast agents were injected intravenously into xenograft mice, following which these mice were scanned at predetermined time points. The MRI images and SER values of three groups were summarized in Fig. 5A and B. In pure Gd-DTPA group, the signal Table 2 Cellular uptake of paramagnetic contrast agents by HUVEC and A549 cells. Paramagnetic liposomes were prepared and added to the cell pellets (n = 4). Compared to pure Gd-DTPA and non-targeted paramagnetic liposomes (NT-Gd-LP), RGD-targeted paramagnetic liposomes (T-Gd-LP) showed greater cells binding affinity. Each data represents the mean ± standard deviation by mmol/l.

HUVEC A549 * #

Gd-DTPA

NT-Gd-LP

T-Gd-LP

0.2937 ± 0.0200 0.1965 ± 0.0200

0.4473 ± 0.0139 0.2375 ± 0.0840

0.5690 ± 0.0410* , # 0.2920 ± 0.0562* , #

p < 0.05 versus pure Gd-DTPA. p < 0.05 versus NT-Gd-LP.

intensity of tumor increased rapidly and reached the maximum after 1 h followed by a speedy decrease. On the other hand, in nontargeted liposomes group, a minor increase in signal enhancement of the tumor within 2 h was observed. Subsequently, the SER value declined and returned to the baseline at 6 h postinjection. By comparison, mice that received injection of RGD-targeted liposomes showed a greater increase in SER of tumor. The maximal SER was found 4 h postinjection which was also the highest in all groups and it maintained a high level till 6 h (p values were 0.05 for 4 h and 6 h, respectively).

3.5. Biodistribution and tumor uptake of Gd The Gd concentration of tumor tissue and other organs were determined to study the distribution and tumor uptake of paramagnetic contrast agents. As shown in Fig. 6, in RGD-targeted liposomes group, the percentage of total Gd injected in tumor tissues increased gradually from (3.71 ± 0.53%) at 30 min to a maximum of (16.27 ± 1.66%) at 4 h postinjection, followed by a slow decrease to (9.11 ± 1.12%) within 24 h. The change of Gd concentration in non-targeted liposomes group was more stable but remained at a low level. On the other hand, in contrast to liposomal contrast agents, pure Gd-DTPA was delivered to tumor tissue and eliminated more rapidly. Gd level of tumor in targeted liposomes group was much higher than pure Gd-DTPA and non-targeted lipo-

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Fig. 5. (A) MRI images of A549 tumor-bearing mice. Mice received (A) Pure Gd-DTPA, (B) RGD-targeted paramagnetic liposomes and (C) non-targeted paramagnetic liposomes injection via tail vein. The tumor regions were circled by dotted lines. (B) Analysis of MRI signal enhancement rate in A549 tumor-bearing mice. Pure Gd-DTPA, non-targeted paramagnetic liposomes (NT-Gd-LP) and RGD-targeted paramagnetic liposomes (T-Gd-LP) were injected via tail vein after which mice were scanned by MRI at predetermined time point (n = 4–5).

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Fig. 6. The biodistribution of Gd-DTPA in A549 tumor-bearing mice (n = 4). (A) The distribution of pure Gd-DTPA, non-targeted (NT-Gd-LP) and RGD-targeted (T-Gd-LP) paramagnetic liposomes in tumor tissues. (B) The distribution of non-targeted paramagnetic liposomes (black column), RGD-targeted paramagnetic liposomes (gray column) group and pure Gd-DTPA (white column) in other organs at 4 h postinjection. (C) Comparison of Gd levels between targeted liposomes (black column) and non-targeted liposomes (white column) at 24 h postinjection. There was more Gd accumulated in tumor tissue but less in blood for targeted liposomes than nontargeted liposomes. *p < 0.05, **p < 0.01 versus non-targeted liposomes.

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somes groups at 2 h, 4 h and 24 h postinjection (p < 0.05). These data clearly illustrated that the RGD-targeted paramagnetic liposomes can efficiently increase the uptake of the Gd-DTPA by tumor tissue. The biodistribution of Gd in blood, liver, spleen and kidney of tumor-bearing mice was also evaluated. At 4 h postinjection, the blood level of liposomal groups was higher than pure GdDTPA group (p values were <0.05, respectively). Furthermore, tissue levels of Gd for liposomal contrast agents were high in RES (reticuloendothelial system) sites (liver, spleen) as well as high Gd level in renal system (kidney) for pure Gd-DTPA. In addition, accumulation of Gd in other organs including heart, lung and muscle was respectively less. This could be explained by different pharmacokinetic characteristics of pure Gd-DTPA and PEG-coating liposomes [21]. At 24 h post injection, the Gd concentration of liver and spleen were still higher than other organs in both liposomal groups. However in targeted liposomes group, there was more Gd accumulated in tumor tissue but less in blood than in non-targeted liposome group, indicating that the specific binding of targeted liposomes to tumor in vivo with a prolonged circulation time.

4. Discussion Molecular imaging, defined as the visualization, characterization and measurement of biological processes at the molecular and cellular levels in humans and other living systems [22], has undergone explosive growth over the last few decades. Through the application of molecular imaging, physiological and pathological information could be acquired by nuclear magnetic resonance and optical imaging technologies. Therefore, molecular imaging is widely applied for the early detection of tumor which is associated with a higher survival rate. There are several commonly used noninvasive techniques for molecular imaging, such as MRI [2], nuclear methods (SPECT and PET) [23], optical fluorescence [24] and ultrasound [25]. As compared to traditional imaging modalities, MRI not only represents the anatomic information of lesions but also has the ability to visualize the physiological and pathological process. In addition, the advantage of high spatial resolution made MRI one of the leading approaches among the above molecular imaging modalities [26]. However MRI suffers from a relatively low sensitivity and the limited specificity. To overcome these disadvantages, contrast agents are used to increase the signal intensity [26]. For clinical diagnosis, paramagnetic contrast agents based on Gd are the most commonly used, especially Gd-DTPA. For sufficient contrast to image sparse molecular biomarkers of tumor cells which is only 10−9 to 10−13 mol/g [27] in vivo, a gadolinium contrast agents concentration of 10−7 mol/g in a tissue is required [28]. It is much higher than the low expression of the biomarkers. In order to solve this problem, one effective approach is to entrap a high payload of contrast agents into nanoparticles. These nanoparticles should deliver enough contrast agents to a tumor tissue, by which the local signal intensity of MRI would be increased [29]. Many forms of nanoparticles are available, such as lipid-based nanoparticles (liposomes [30,31] and solid lipid nanoparticles [32,33]), polymeric nanoparticles [34,35] and carbon nanotubes [36,37]. Due to the relatively large aqueous core with a high drug-to-carrier ratio and no toxicity in prescription dose, liposomes composed of amphiphilic lipid bilayers are the most intensively studied as carriers of MRI contrast agents. Furthermore, incorporation of PEG into lipid bilayers could effectively “sterically stabilize” liposomes so as to increase its half life by interfering the interaction between liposomes and approaching proteins, neighboring cells or other liposomes [38,39]. To specifically target cellular biomarkers of tumors, many varieties of ligands can potentially be conjugated to the lipo-

somes surface, including antibodies [18,40,41], peptides [2,42] and polysaccharides [43,44]. The RGD-motif/integrin system is currently intensively investigated in angiogenesis targeted tumor studies. RGD-containing peptides which preferentially bind to ␣␯␤3 or other integrins have been widely used for tumor targeted drug delivery [13,14,45]. However, there are very few studies of RGD-peptides based liposomal imaging agents [2,13,19,46]. In our study, we selected a cyclic RGDcontaining peptide RGD10 (DGARYCRGDCFDG) which was isolated from phage display libraries. The RGD10 peptide has similar high affinity for ␣␯␤3 integrin as RGD4C peptide which was one of the most investigated peptide to date, but unlike the RGD4C peptide it does not fold into other different cyclic structures because RGD10 has only one disulfide bond [45,47]. Additional study also illustrated the specific binding affinity of RGD10-conjugated liposomes to integrin-high expressed cells. A variety of methods could be used to conjugate RGD-peptides to the PEG-coating liposomes surface. One of the main structures is formed by coupling the peptides to the phosphatidylcholine headgroups directly. However, due to interference by the PEG polymer chains, inhibition of the specific binding of the peptides to tumor cells occurred [48]. To solve this problem, another structure was constructed by attaching the peptide to the free distal end of PEG polymer chain in many recent investigations. This successfully increased the binding affinity of liposomes to tumor cells. Another alternative approach is incorporation of lipopeptide in the lipid bilayer of targeted liposomes [49–51]. In our study, we synthesized a novel lipopeptide which contained a RGD peptide headgroup and palmitic acid anchor. Moreover, we attached the RGD peptide via a KGG spacer to the anchor to minimize the impact of the steric polymer chains. The tripeptide KGG was firstly used to be the headgroup of the lipophilic peptides for gene delivery and demonstrated to be effective [52]. Additionally, the lysine region possesses two amino which could binds to single palmitic acids chains separately to mimic a phospholipids structure and firmly incorporates the lipopeptide into the lipid bilayer structure (Fig. 1) [53]. Using fluorescent microscopy, we demonstrated the specific binding of targeted liposomes to HUVEC or A549 cells. Besides, the binding activity was determined by fluorescence signal intensity analysis. It was illustrated that the binding affinity of targeted liposomes was much higher than non-targeted liposomes. Additionally, the competition experiments demonstrated that the binding was mediated via RGD motif/integrin interaction. Paramagnetic liposomes were also incubated with cells to further confirm in vitro targeting of liposomes. As expected, the cellular uptake of Gd in targeted liposome group was higher than non-targeted or pure Gd-DTPA group, corresponding to a higher accumulation of RGDtargeted liposomes into the cells. For in vivo MRI study, a human lung cancer xenograft model was established. In order to investigate the ability for early diagnosis, the volumes of tumors were limited to 50–100 mm3 when used for MRI scan. The images showed that on the 1.5 T MR scanner the tumor was enhanced clearly by targeted paramagnetic liposomes. Moreover, analysis of the MRI signal intensity revealed different time-dependant enhancement patterns of all three groups. Pure Gd-DTPA showed a rapid clearance from tumor tissue with mild signal enhancement because the water-soluble Gd chelates diffuses from the tumor tissue easily. In comparison to paramagnetic liposomes, the SER value of tumor was higher. The signal enhancement of tumor decreased after peaking within 2 h postinjection in non-targeted group compared to a continued rise till the highest value was attained in targeted liposomes group. We suggest that the accumulation of non-targeted paramagnetic liposomes in tumor is caused by only extravasations of the liposomes but not by specific binding to the tumor tissue. In contrast, because of the specific binding to neovessel endothelial cells and tumor cells, theo-

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retically the targeted paramagnetic liposomes in blood circulation could gradually accumulated in the tumor tissue to enhance the signal intensity. In addition, with concern to the biodistribution study, we found that in non-targeted liposomes group there was no significant difference of Gd level between tumor and control muscle tissues. Thereby the MRI signal enhancement rate which also compared the signal intensity discrimination of tumor and muscle tissues was low. But in targeted liposomes group, the Gd concentration of tumor tissue was higher than that of muscle, which resulted in substantial signal enhancement. 5. Conclusion In conclusion, we designed and synthesized a novel lipopeptide which was constructed by conjugation of a RGD-peptide headgroup to palmitic acid anchor via a KGG space. Fluorescent and paramagnetic liposomes were prepared by thin film hydration methods with RGD lipopeptides incorporated into lipid bilayers for specific targeting to endothelial cells and tumor cells in vitro. Fluorescent microscopy and spectrophotometer test illustrated the specific binding activity of RGD-targeted liposomes. The same results were obtained in cellular uptake study of paramagnetic liposomes. In vivo MRI scanning also demonstrated that the signal enhancement in tumor after injection of RGD-targeted liposomes was significantly higher than non-targeted liposomes and pure Gd-DTPA. Thus, we expect that our RGD-targeted paramagnetic liposomes can be used to deliver a sufficient amount of contrast agents into tumor and may provide an effective means for noninvasive diagnosis of early tumor. Conflict of interest statement We declare that we have no financial and personal relationships with other people or organizations that can inappropriately influence our work, there is no professional or other personal interest of any nature or kind in any product, service and/or company that could be construed as influencing the position presented in the manuscript. Acknowledgment This work was financially supported by a grant from National High Technology and Development Program of the People’s Republic of China (2008AA02Z442). References [1] Daldrup H, Shames DM, Wendland M, et al. Correlation of dynamic contrast-enhanced MR imaging with histologic tumor grade: comparison of macromolecular and small-molecular contrast media. AJR Am J Roentgenol 1998;171(4):941–9. [2] Mulder WJ, Strijkers GJ, Habets JW, et al. MR molecular imaging and fluorescence microscopy for identification of activated tumor endothelium using a bimodal lipidic nanoparticle. Faseb J 2005;19(14):2008–10. [3] Kinoshita M, Yoshioka Y, Okita Y, Hashimoto N, Yoshimine T. MR molecular imaging of HER-2 in a murine tumor xenograft by SPIO labeling of anti-HER-2 affibody. Contrast Media Mol Imaging 2010;5(1):18–22. [4] Weissleder R, Simonova M, Bogdanova A, Bredow S, Enochs WS, Bogdanov AJ. MR imaging and scintigraphy of gene expression through melanin induction. Radiology 1997;204(2):425–9. [5] Burnett CA, Xie J, Quijano J, et al. Synthesis, in vitro, and in vivo characterization of an integrin alpha(v)beta(3)-targeted molecular probe for optical imaging of tumor. Bioorg Med Chem 2005;13(11):3763–71. [6] Folkman J. Tumor angiogenesis: therapeutic implications. N Engl J Med 1971;285(21):1182–6. [7] Folkman J. Angiogenesis in cancer, vascular, rheumatoid and other disease. Nat Med 1995;1(1):27–31. [8] Hood JD, Cheresh DA. Role of integrins in cell invasion and migration. Nat Rev Cancer 2002;2(2):91–100.

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