Gadolinium-loaded polychelating amphiphilic polymer as an enhanced MRI contrast agent for human multiple myeloma and non Hodgkin's lymphoma (human Burkitt's lymphoma)

Abstract

Liposomes, loaded with gadolinium (Gd) ions using different membrane-incorporated chelating lipids and functionalized with monoclonal anti-CD138 (syndecan-1) antibody were prepared. Nuclear Magnetic Resonance Dispersion (NMRD) analysis showed that use of the polychelating amphiphilic polymer (PAP) increases both the Gd content and the spin–lattice relaxivity of the Gd-loaded-PAP–liposomes as compared to Gd–DTPA–BSA equivalents. The potential application of contrast syndecan-1– and Rituximab–liposomes, for application as a novel minimally invasive diagnostic agent for multiple myeloma and non Hodgkin's lymphoma was investigated.

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Introduction

Magnetic resonance imaging (MR) has become one of the leading non-invasive modalities used clinically. It is widely applied to obtain images of pathological tissue, as it is capable of generating three-dimensional non-invasive images of opaque tissues at near cellular resolution and may have the greatest potential to exploit the possibilities that molecular imaging presents. MRI exhibits good spatial resolution but low sensitivity in the region of 10⁻³ to 10⁻⁵ mol L⁻¹. This low sensitivity can be overcome by using contrast agents (CA) with high relaxivity. Although MRI is a non-invasive technique around 25% of the current MR-examinations include a contrast agent. These agents consist of molecules which incorporate a paramagnetic metal ion, most commonly gadolinium or iron. The inclusion of a paramagnetic agent, such as Gd, increases the R₁ and R₂ relaxation rates. However, the best visualization is obtained using T₁-weighted images because in tissue the percentage change in R₁ is much greater than in R₂. Gd is an example of paramagnetic agent which directly affects water protons in its close vicinity. Hence, the influence of paramagnetic agents is very local and they should ideally be in contact with water with adequate exchange. In contrast, superparamagnetic agents (i.e. Fe) affect the magnetic field independently of their environment and, thus, their influence in terms of contrast extends well beyond their immediate surroundings.¹

To enhance the MR-efficacy and/or to improve the site-specific delivery of gadolinium (Gd) based MR-contrast agents, various systems such as nanoparticles have been suggested as carriers.² Nanoparticles, including liposomes and other nanoscale constructs (solid lipid nanoparticles, polymeric nanoparticles, nanocrystals, dendrimers and carbon nanotubes) are the focus of intense research, due to potential applications in the biomedical, optical, and electronic fields.³ In this research liposomes were generated to target human multiple myeloma and human Burkitt's lymphoma for diagnostic applications.

Human multiple myeloma (MM) is a plasma B cell malignancy, which occurs in bone marrow. MM accounts for slightly more than 10% of all haematological malignancies.⁴ CD138 (syndecan-1), is over-expressed on the major types of human multiple myeloma cancer cells in bone marrow and is a good marker of this malignancy.⁵

Lymphoma is a broad term that encompasses a wide group of malignancies of the lymphoproliferative system of different biology and prognosis. Lymphoma is generally subdivided into two main groups, Hodgkins Lymphoma (HL) and Non-Hodgkins Lymphoma (NHL). NHL is the most common primary haematological malignancy and its incidence is on the rise. Greater than 90% of all NHL are of B-cell origin. At present surgical excision of a lymph node followed by histological and immunohistochemical examination of the lymph node is the gold standard for diagnosing lymphoma. CD20 is an antigen expressed specifically by B cells, and is homogeneously expressed on more than 90% of B cell lymphomas at a density of 50 000 to 200 000 molecules per lymphoma cell.⁶

Targeting of MRI contrast agents to distinct molecules associated with tissue pathology could enable more specific diagnosis by MRI leading to characterisation of the disease at the molecular level in vivo. Qiao et al.⁷ demonstrated that a Gd³⁺-chelated protein-based (ProCA1) MRI agent, which was HER2-targeted antibody-conjugated, showed higher relaxivity than the non-targeted equivalent in ovarian (SKOV-3) and breast (MDA-MB-231) cancer cell lines in vivo. Furthermore, Wang et al.⁸ prepared novel anti-BRCAA1 (a protein over-expressed in 64% of gastric cancer tissue) monoclonal antibodies that were linked to fluorescent magnetic nanoparticles (FMNPs) composed of silica-coated quantum dots and super-paramagnetic nanoparticles. They were effective for in vivo dual modal imaging by fluorescence and magnetic resonance.

Antibody-labelled composition-optimised liposome-encapsulated gadolinium have significantly improved kinetics due to longer blood circulation, selective targeting/distribution to tumors relative to healthy tissues, improved extravasation from blood at the tumor's location and enhanced tumour binding due to specific antibody docking.⁹

Targeted Gd-loaded nanoparticles can overcome specificity problems related to non-targeted paramagnetic agents. Moreover, an amplification strategy for gadolinium-containing contrast agents is required to enhance sensitivity. Gadolinium-containing contrast agents currently available require a tissue concentration in the order of 10⁻⁷ mol g⁻¹ to obtain sufficient contrast in the resulting image.^1,10 This level is too high to facilitate imaging of sparse molecular biomarkers expressed on cells since such molecules are normally present in low concentrations of 10⁻⁹ to 10⁻¹³ mol g⁻¹.¹¹ One solution is to use nanoparticles since a high payload of contrast agent can be incorporated per single particle, thus increasing the effective relaxivity per particle, and, consequently, the signal detected during imaging increases.¹² PAP (polychelating amphiphilic polymer) chelates many contrast agents per molecule and is capable of increasing relaxivity per particle. PAP was previously synthesized and used for tumor-targeting of PEGylated 2C5-conjugated-liposomes in vitro¹³ and in vivo.¹⁴

The long term goal of this research was the generation of a novel contrast agent for minimally invasive diagnosis of both human multiple myeloma and Burkitt's lymphoma in humans. In the reported study, Gd-loaded PAP-containing PEGylated-liposomes were evaluated as potential tumor-targeted contrast agents for MRI. The T₁ relaxation time of these liposomes was compared with Gd–DTPA–BSA-loaded liposomes and it was demonstrated that Gd-loaded-PAP–liposomes show higher r₁ relaxivity than Gd–DTPA–BSA (diethylenetriaminepenta-acetic acid α,ω-bis(8-stearoylamido-2,6-dioxaoctylamide)gadolinium salt)-liposomes loaded with monomeric Gd–DTPA, at the same molar ratio as the chelate. It was also shown that targeted anti-CD138-conjugated-Gd–PAP–liposomes bound significantly more to human multiple myeloma cell lines compared to plain contrast liposomes (blank liposomes without antibody) or negative control antibody-conjugated-Gd–PAP–liposomes. These results can be considered as an important step in the development of targeted contrast agents for minimally invasive diagnosis of MM and Burkitt's lymphoma.

The paramagnetic gadolinium, Gd³⁺, ion has been long used to overcome sensitivity issues related to MR imaging as its presence shortens the T₁ and T₂ relaxation times of the surrounding water.¹⁹ It has been found that liposome-based contrast agents, including Gd ions, demonstrate potentially promising results for use as diagnostic imaging agents for different types of cancer.^{14,17,20–22} The efficacy of the Gd³⁺ – bearing species for altering the relaxation properties of the water ¹H nuclei can be quantified by the relaxivities, r₁ or r₂, which are defined as the change in R₁ or R₂ relaxation rates, per unit concentration of the contrast agent; r₁ and r₂ have units of s⁻¹ mM⁻¹.

The traditional method of encapsulating mainly water soluble contrast agents, used Gd³⁺ chelates, within the liposome lumen and has the disadvantage that the entrapped agents easily ‘leak out’ before the agent can reach the target site. This problem can be avoided by chelation of Gd³⁺ with lipid-compatible molecules like DTPA–BSA, DOTA–DSA or NGPE–PLL–DTPA that readily insert their chains into the membrane bilayer. In this way, Gd³⁺ can be anchored to the liposome surface instead of being encapsulated in the lumen. Kamaly et al. found that a liposomal formulation containing Gd–DOTA–DSA was non-toxic to HeLa cells, as demonstrated by the LDH viability assay.²³ Hence we prepared liposomes using Gd-loaded NGPE–PLL–DTPA (PAP) and Gd-loaded DTPA–BSA, to anchor the contrast agent to the liposome bilayer whilst hopefully avoiding toxic effects on cells. These complexes are particularly advantageous as a high payload of ions can be incorporated in a single liposome, thus increasing the effective relaxation rate enhancement achievable.²⁴ Previously Torchilin et al. showed that the NGPE–PLL-CBZ complex generates up to 11 amine groups that are available to undergo reaction with DTPA, which can subsequently be loaded with Gd ions²⁵ with a resulting average loading of 6 ions. Erdogan et al. (2006) demonstrated that liposomes loaded with Gd via membrane-incorporated polychelating amphiphilic polymer (PAP) significantly increased the Gd content and relaxation rate, R₁, of PEGylated liposomes.¹³

Materials and methods

Materials

L-α-Phosphatidylcholine, hydrogenated (egg, chicken) (egg PC), cholesterol, 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-[carboxy(polyethylene glycol)2000]-carboxylic acid (PEG–DSPE–carboxylic acid), 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-[methoxy(polyethylene glycol)-2000] (DSPE–PEG-methoxy), 1,2-dioleoyl-sn-glycero-3-phosphoethanolamine-N-(glutaryl) (sodium salt) (NGPE), 1,2-dioleoyl-sn-glycero-3-phosphoethanolamine-N-(carboxyfluorescein) (FITC–PE), polycarbonate membranes (‘cut-off’ 100 nm) were purchased from Avanti Polar Lipids, Inc. (Delfzyl, Netherlands). Poly-ε-CBZ-DL-lysine (MW 5500) (PLL-CBZ), N,N′-dicyclohexyldiimide (DCC), N-hydroxysuccinimide (NHS), GdCl₃·6H₂O, succinic anhydride, dimethylformamide (DMF) and penicillin–streptomycin were obtained from Sigma-Aldrich (Hannover, Germany). Vivaspin columns (5000 Da) were supplied by Sartorius (Goettingen, Germany).

Chemicals used to prepare buffer solutions were analytical grade.

RPMI 1640 medium and foetal bovine serum (FBS) were obtained from Bio-sciences (Crofton Road, Dun Laoghaire, Co. Dublin, Ireland).

Diethylenetriaminepentaacetic acid α,ω-bis(8-stearoylamido-3,6-dioxaoctylamide)gadolinium salt (Gd–DTPA–BSA monohydrate) was purchased from Azopharma (St. Louis, MO, USA).

U266 and RPMI8226 human multiple myeloma were supplied by Deutsche Sammlung von Mikroorganismen und Zellkulturen (DSMZ, Braunschweig, Germany). The RAMOS, L428 Hodgkin's disease-derived cell line was kindly donated by Dr Dermot Walls, School of Biotechnology, Dublin City University, Ireland.

Methods

Polychelating amphiphilic polymer (PAP–DTPA–polylysyl-N-glutaryl-phosphatidylethanolamine) synthesis. Polychelating Amphiphilic Polymer (PAP) was synthesized as described previously.^15,16 Briefly, 1 mL of 25 mg mL⁻¹ NGPE in chloroform (1,2-dioleoyl-sn-glycero-phosphoethanolamine-N-(glutaryl)) was activated with 0.5 mL of N-(3-dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride (20 mg) in the presence of 0.5 mL of N-hydroxysuccinimide (13 mg) for 2–3 hours at room temperature. Subsequently, 40 μL of triethylamine was added to make the solution basic to allow conjugation with poly-ε-CBZ-DL-lysine (PLL-CBZ).

All NH₂ groups of PLL-CBZ were protected by a CBZ group, except one which was linked to the activated carboxylic group of NGPE. To perform this reaction, 186 mg of PLL-CBZ was dissolved in 4 mL of DMF solvent and pooled together with activated NGPE phospholipids. The reaction was performed for 12 hours at room temperature with stirring. Chloroform was then evaporated from the mixture and product was precipitated using distilled water. The NGPE–PLL-CBZ product was freeze-dried for 48 hours and 193 mg of the NGPE–PLL-CBZ compound obtained. The NGPE–PLL-CBZ complex was de-protected from the CBZ group using 2 mL of 33% (v/v) hydrogen bromide in acetic acid for 2 hours at room temperature with stirring. The complex was precipitated using diethyl ether and freeze-dried for 12 hours. De-protected complex of NGPE–PLL (107 mg) was obtained. De-protection of the NGPE–PLL-CBZ complex generated a high number of NH₂ groups which were available for reaction with DTPA. NGPE–PLL was treated with DTPA anhydride in the presence of triethylamine for 16 hours at room temperature with stirring. Succinic anhydride was added and the reaction was allowed to proceed for 1 hour at room temperature. The complex was then dialysed against distilled water (‘cut-off’ 3500 kDa) to remove unbound succinic anhydride (Fig. 1).

Fig. 1 Polychelating amphiphilic polymer (PAP) synthesis. Poly(CBZ)Lys was conjugated with NGPE (1,2-dioleoyl-sn-glycero-phosphoethanolamine-N-(glutaryl)) generating poly(CBZ)Lys–NGPE-conjugated product. Then CBZ was de-protected from poly(CBZ)Lys–NGPE. The de-protected compound was then linked with DTPA cyclic anhydride generating NGPE–polylysine–DTPA (NGPE–PLL–DTPA) compound, called polychelating amphiphilic polymer (PAP). PAP bears many DTPA ligands ready for metal chelation.

Loading of chelating polymers with Gd ions. GdCl₃·6H₂O was added to DTPA–PLL–NGPE suspended in 2 mL of pyridine and stirred for 4 hours at room temperature. The reaction mixture was dialyzed for 24 hours against deionized water and vacuum dried. The procedure gave a polymer preparation containing ≥30% (weight) of gadolinium.¹⁶

Preparation of Gd-loaded-DTPA–BSA- and Gd-loaded-PAP-containing PEGylated liposomes. Uniform mixtures of 63.25 mol% egg PC, 30 mol% cholesterol, 5 mol% PEG₂₀₀₀–PE and 1.75 mol% Gd–PAP or 1.75 mol% Gd–DTPA–BSA and cholesterol were prepared by dissolving the lipids in CHCl₃. The solvent was removed under a slow stream of N₂ over 30 minutes. Liposomes were prepared by hydrating the mixtures in 2.2 mL of distilled water for Gd–DTPA–BSA–liposomes and 0.7 mL of distilled water for Gd–PAP–liposomes. The suspensions were then passed through an Avanti Polar Lipids mini extruder (Alabaster, AL) containing polycarbonate membranes with a pore size of 0.1 μm (Whatman Nuclepore; Clifton, NJ). Each suspension was passed through the mini extruder 10 times.

Particle size measurements. The average hydrodynamic size of liposomes and anti-CD138-conjugated-liposomes was determined by photon correlation spectroscopy (PCS), using a NanoZS (Malvern Instruments, Malvern UK). The instrument uses a detection angle of 173° and a 3 mW He–Ne laser operating at a wavelength of 633 nm. The mean hydrodynamic diameter is based upon the intensity of scattered light. Photon Correlation Spectroscopy (PCS) can also provide information about the zeta potential of liposomes. The zeta potential is important in the assessment of liposome stability and propensity for aggregation. Zeta potential indicates the degree of repulsion between adjacent, similarly charged particles in dispersion. For molecules and particles that are small enough, a high zeta potential will confer stability, e.g., the solution or dispersion will resist aggregation. When the potential is low, attraction exceeds repulsion and the dispersion will break and flocculate. Colloids with high zeta potential (negative or positive) are electrically stabilized while colloids with low zeta potentials tend to coagulate or flocculate.³³

Relaxivity of Gd-loaded-DTPA–BSA- and Gd-loaded-PAP–liposomes. NMRD profiles were recorded to characterize the magnetic resonance characteristics of the contrast liposome samples. The relaxometry experiments were performed on a Spinmaster FFC-2000 Fast Field-Cycling NMR Relaxometer (Stelar, Mede, Italy) using a sample volume of 0.5–1 mL. Measurements were carried out at 25 °C and controlled within ±1 °C through the use of a thermostatted airflow system. To measure T₁ at frequencies higher than 20 MHz, a reconditioned Bruker WP80 variable field magnet was used. This allowed measurements of T₁ from 25–70 MHz.

The gadolinium content in contrast liposomes (Gd–DTPA–BSA–liposomes and Gd–PAP–liposomes) was then determined using Inductively Coupled Plasma-Atomic Emission Spectroscopy (ICP-AES). For ICP-AES analyses, liposomes were dissolved in 0.2% (v/v) Triton X-100 and made up quantitatively for analysis. A linear calibration curve, with R² > 0.99, for GdCl₃ standards in the same concentration ranges was established, and the gadolinium content of the liposome suspensions determined. Spin–lattice relaxivity, r₁, values were then obtained by normalizing the measured rates for the gadolinium content, according to:

where R_1,obs is the measured relaxation rate of the liposome suspension, R_1,H2O is the relaxation rate of water (no liposomes present) and [Gd] is the concentration of gadolinium contrast agent, in mM.

Cell lines used. Human multiple myeloma (U266, RPMI8226), Hodgkin's disease-derived (L428) and human Burkitt's lymphoma (RAMOS) cell lines were cultured in RPMI1640 supplemented with 10% (v/v) and 1% (v/v) penicillin–streptomycin at 37 °C in a humid 5% (v/v) CO₂ atmosphere.

Preparation of anti-CD138–Gd–PAP-conjugated-liposomes. Anti-CD138 monoclonal antibody was covalently bound to DSPE–PEG–COOH (1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-[carboxy(polyethylene glycol)2000]-carboxylic acid) lipid. An amide bond was formed in two stages to achieve conjugation. In the first stage, a free carboxylic group on the terminal end of the PEG chain was activated with water-soluble N,N′-dicyclohexyldiimide (DCC) and N-hydroxysuccinimide (NHS). In the second stage, antibody was resuspended in 50 mM K₂HPO₄ buffer, pH 8.0, using a Vivaspin column (‘cut-off’ 5000 Da). The activated carboxyl groups reacted with amino groups on the antibody to create an amide bond. The presence of esterified DSPE–PEG–COOH lipid was confirmed using thin layer chromatography (TLC). For TLC both DSPE–PEG–COOH and esterified DSPE–PEG–COOH were run in chloroform–methanol (2 : 1) mobile phase on a silica gel plate (60 Å pore diameter). The dried silica plate was placed in an iodine tank for visualization of the separated lipid spots.

Anti-CD138 (syndecan-1) monoclonal antibody (0.2 mg) was at a 40 molar excess of esterified DSPE–PEG–COOH lipid generating DSPE–PEG–CO–NH–anti-CD138-antibody. A co-incubation method was used to prepare anti-CD138-conjugated liposomes.^13,14,17,18

To prepare contrast immunoliposomes, 62.25 mol% egg PC, 30 mol% cholesterol, 5 mol% PEG₂₀₀₀–PE, 0.5 mol% FITC–PE and 1.75 mol% Gd–PAP were used. Solvent was evaporated under nitrogen. The lipid film was hydrated in 1 mL 0.1 M PBS, pH 7.3, by vortexing for 25 min at room temperature. Liposomes were then extruded through a 100 nm polycarbonate membrane. Then, 2 mol% of DSPE–PEG–CO–NH–anti-CD138-antibody previously prepared was co-incubated with liposomes by stirring overnight at 4 °C.

Plain Gd–PAP–liposomes (liposomes without antibody conjugated) and Herceptin®–Gd–PAP–liposomes were prepared as a negative control. Herceptin® (Trastuzumab) is a humanized monoclonal antibody which interferes with HER2 receptor overexpressed in breast cancer cells.

Separation of non-conjugated anti-CD138 antibodies from contrast anti-CD138-conjugated-liposomes was performed by dialysis against 0.1 M PBS, pH 7.3, for 2 days (‘cut-off’ 300 000 Da).

Preparation and purification of fluorescently-labelled Rituximab-conjugated-liposomes. Briefly, 64.5 mol% of egg PC, 30 mol% of cholesterol, 3 mol% of DSPE–PEG–COOH and 0.5 mol% of rhodamine–PE were used to prepare a lipid film. The lipid film was hydrated using 1 mL of 0.1 M PBS, pH 7.4 and vortexed. It was then extruded through a 100 nm polycarbonate membrane for liposome generation.

The Rituximab antibody (0.6 mg) was conjugated to 2 mol% of DSPE–PEG–COOH by firstly esterifying the DSPE–PEG–COOH using DCC and NHS. Rituximab is a chimeric monoclonal antibody that binds to CD20 which is primarily found on B cells.

A forty molar excess of this esterified lipid was conjugated to the Rituximab monoclonal antibody by incubation in 0.1 M PBS, pH 8.5, overnight at 4 °C. This Rituximab-conjugated-phospholipid solution was dialyzed using 3500 Da ‘cut-off’ dialysis tubing against 0.1 M PBS, pH 7.4, for 2 h to remove DMSO, in which esterified lipid was resuspended. Subsequently the mixture was co-incubated with the previously prepared liposomes at 4 °C overnight.

Fluorescently-labelled-liposomes were dialyzed against PBS, pH 7.4, (‘cut-off’ 300 000 Da) for 2 days at 4 °C to remove unbound antibody.

In vitro cell binding experiments. Two human multiple myeloma (U266; RPMI8226) and the L428 Hodgkin's disease-derived cell lines were tested to analyze CD138 protein expression levels. Cells (1 × 10⁶) were incubated with 0.25 μg mL⁻¹ of FITC-labelled anti-CD138 antibody. Non-specific binding of antibodies was assessed by labelling cells with an isotype control antibody (FITC-labelled κ isotype mouse IgG). Samples were well mixed and incubated for 30 min at 4 °C. Finally, the cells were washed three times by centrifugation at 3000g for 5 min using Z 233 MK-2 centrifuge (HERMLE Labortechnik, UT 84037, USA) and resuspended in 200 μL 4% (v/v) formaldehyde/PBS. Samples were transferred to FACS tubes and were analyzed immediately. Antibody-labelled cells were identified and analysed for surface fluorescence at 488/530 nm on a FACSscan flow-cytometer (Becton-Dickinson). Fluorescence data were collected and analysed using CellQuest software (Becton Dickinson).

Binding of the Rituximab monoclonal antibody to RAMOS (human Burkitt's lymphoma) and L428 (Hodgkin's disease-derived) cells was also tested by flow cytometry. RAMOS and L428 cells (10⁶) were incubated with 1 μg mL⁻¹ of Rituximab or Avastin® antibody for 30 min at 4 °C. The cells were then washed with PBS, pH 7.4, and incubated with a 1 : 2000 dilution of Cy5-labelled anti-human IgG secondary antibody for 30 min at 4 °C. Cells were washed three times with FACS buffer (0.1 M PBS, pH 7.4 with 2% (v/v) FBS and 0.05% (w/v) NaN₃) and immediately analyzed on a FACS Calibur instrument.

Magnetic resonance imaging. Contrast liposome, anti-CD138–Gd–DTPA–BSA-conjugated-liposome or anti-CD138–Gd–PAP-conjugated-liposome cell interactions were visualized using a 1.5 T scanner (Gyroscan Intera; Philips, Netherlands) at the Radiology Department, Cappagh National Orthopaedic Hospital, Finglas, Dublin 11, Ireland. Samples were placed into a container filled with water. The software used was Release 8.0 with SENSE. T₁-weighted images were performed using spin-echo (SE) imaging sequence (repetition time, TR = 15 ms; echo delay time, TE = 100 ms; matrix 256 × 256).

Interaction of fluorescently-labelled-anti-CD138–Gd–PAP-conjugated-liposomes with human multiple myeloma cancer cell lines. The interactions of fluorescently-labelled contrast anti-CD138–Gd–PAP-conjugated-liposomes with cells were studied. The anti-CD138-conjugated-liposomes were prepared from 62.75 mol% egg PC, 30 mol% cholesterol, 5 mol% PEG₂₀₀₀–PE, 0.5 mol% FITC–PE and 1.75 mol% Gd–PAP. Blank liposomes (liposomes without antibody attached) were also prepared as a negative control. Approximately 5 × 10⁵ cells (U266, RPMI8226 of multiple myeloma and L428 of Hodgkin's disease-derived cancer cell lines) were incubated with 0.1 mg mL⁻¹ of liposomes in ‘serum-free’ medium for 1–1.5 hours at 37 °C in a 5% (v/v) CO₂ humidified atmosphere. Cells were then washed three times with 0.1 M PBS, pH 7.3, and finally resuspended in FACS buffer (0.1 M PBS, pH 7.4 with 2% (v/v) FBS and 0.05% (w/v) NaN₃) and immediately analysed by FACSscan flow-cytometry (Becton-Dickinson). Immunoliposome-labelled cells were identified and analysed for FITC fluorescence at 488/530 nm. Fluorescence data were collected and analysed using CellQuest software (Becton Dickinson).

Results and discussion

Synthesis and physical characterisation

The synthetic scheme for the preparation of Gd-loaded polychelating amphiphilic polymer (PAP) is shown in Fig. 1. This lipid provides the possibility of increasing the Gd-loading, as compared to lipidic carriers such as Gd–DTPA–BSA, which can accommodate only a single ion. In the current study, the mole fraction of the Gd–lipid was fixed at 1.75 mol%, for both types of lipid, as the report by Weissig et al. (2000) suggested that this is an optimal molar fraction for contrast liposome preparation.¹⁶ The size of contrast liposomes was determined to be 150.3 nm (PDI 0.18) for Gd–DTPA–BSA–liposomes and 120.4 nm (PDI 0.25) for Gd–PAP–liposomes, using PCS. Indeed, PCS shows that liposome suspensions of these two formulations were stable for at least 14 days, and, on exposure to strong magnetic fields. Zeta potential for Gd–DTPA–BSA was approximately −40 mV with variations in the 5–7 mV range. Elemental analysis, by ICP-AES showed a gadolinium concentration of 2.18 mM of in Gd–DTPA–BSA–liposomes and 4.46 mM in Gd–PAP–liposomes. This is equivalent to yields of 67 and 60% for Gd–DTPA–BSA– and Gd–PAP–liposomes, respectively, following all processing steps.

The spin–lattice relaxivity r₁ of Gd–PAP–liposomes was found to be 33 s⁻¹ mM⁻¹ (Gd concentration) compared to 15 s⁻¹ mM⁻¹ for Gd–DTPA–BSA–liposomes, at a measurement field of 60 MHz, which is within the clinical MRI range. Given that the Gd–PAP ligand typically contains 6 ions, this surprising result suggests that a single Gd–PAP ligand could potentially provide improved image contest, under T₁-weighting, by a factor of ∼13 per lipid incorporated.

To further investigate the unusual magnetic resonance characteristics of the contrast liposomes NMRD profiles were recorded for the suspensions at 25 °C, as shown in Fig. 2. The profile for Gd–DTPA–BSA–liposomes is found to be similar to those for related gadolinium-containing suspensions. A low frequency plateau, at 13–14 s⁻¹ mM⁻¹, an r₁ maximum at 15–25 MHz and a decrease in r₁ into the clinical MRI range are very characteristic. These features have been reported in the profiles of several amphiphilic gadolinium chelates included in liposomal formulations. For instance, Gd–DMPE–DTPA with OPPC,²⁶ or Gd–DTPA bisamides with DPPC.²⁷ The profile of Gd–PAP–BSA–liposomes shows a significant increase in relaxivity at all frequencies, but in particular in the clinical range. It should be emphasised that the Gd–PAP– and Gd–DTPA–BSA–liposomes are of the same size and are formulated with the same fraction of Gd-bearing lipid.

Fig. 2 Water ¹H NMRD profiles recorded at 25 °C for solutions of paramagnetic liposomes containing Gd–DTPA–BSA () and Gd–PAP (); the average hydrodynamic sizes of the suspensions were 171 nm (PDI 0.23) and 175 nm (PDI 0.22), respectively. The solid lines are simulations of the NMRD response using the modified SBM model, see text. The inset shows typical linear dependence of R₁ on Gd concentration for Gd–DTPA–BSA liposomes at 10 MHz with an excellent correlation of R² = 0.98.

The NMRD profiles have been fitted only in the high field region (>3 MHz) because of the well-known limitations of the Solomon–Bloembergen–Morgan (SBM) relaxation theory at low magnetic fields strengths under the slow motion regime.²⁸ In order to separate the local molecular rotation of the chelates (characterized by a correlation time τ_RL) from the global tumbling motion of the nanoparticle (τ_RG), the Lipari–Szabo model free approach for the rotational dynamics has been incorporated into SBM theory.²⁹ The best fit was obtained by fixing the number of coordinated water molecules (q = 1), their distance from the paramagnetic metal ion (r_GdH = 3.0 Å), the values for the distance of closest approach of the outer sphere solvent molecules to Gd³⁺ (a = 4.0 Å) and for the relative diffusion coefficient of solute and solvent molecules (D = 2.24 × 10⁻⁵ cm² s⁻¹). The adjustable parameters used were the rate of water exchange (k_ex = 1/τ_M), the parameters describing the electronic relaxation times T_1,2e (Δ², τ_V), τ_RG and τ_RL, and the order parameter S² which represents their degree of correlation (0 < S² < 1). In addition, based on published data in the case of similar DTPA derivatives, the parameter k_ex was only allowed to vary in the range 1–10 × 10⁶ s⁻¹.^28–31 Moreover, the parameters Δ² and τ_V should only be regarded as empirical fitting parameters and do not assume a clear physical meaning for these slowly tumbling systems. The best-fit parameters (Table 1) indicate a pronounced difference between the global motion of the nanoparticle and the local rotation of the complex around the lipophilic chains for Gd–DTPA–BSA–liposomes.

Table 1 Selected parameters obtained from the analysis of the 1/T₁ NMRD profiles (298 K) of paramagnetic liposomes containing Gd–DTPA–BSA and Gd–PAP

Parameter	Gd–DTPA–BSA	Gd–PAP
a At 40 MHz and per Gd. The error bars represent an estimate of the uncertainty in the parameters based on a sensitivity analysis of the fit to the data.
r1p/mM^−1 s^−1^a	17.1 ± 0.2	33.8 ± 0.3
Δ^2/10^19 s^−2	1.9 ± 0.2	2.7 ± 0.2
τV/ps	32 ± 3	9 ± 1
τRL/ns	0.70 ± 0.09	1.7 ± 0.2
τRG/ns	6.6 ± 4	6.6 ± 2
S^2	0.13 ± 0.03	0.33 ± 0.05
^310kex/10^6 s^−1	1.7 ± 0.3	10.0 ± 0.5

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This high degree of rotational flexibility, revealed by the low value of S², is associated with a slow rate of bound water exchange and both of these features represent a limitation to the relaxivity at imaging fields. This scenario, which was previously identified,³² is in line with the physical expectation for DTPA–bisamide derivatives. In this case, the slow rate of water exchange represents the dominant factor that limits the relaxivity at both low and high fields. The effectiveness of the liposomes as MRI probes is significantly enhanced by restricting rotational flexibility of the Gd-chelates incorporated in the membrane and, at the same time, increasing k_ex. The improved motional coupling imposed by the spatial restriction due to the coordination of multiple lanthanide ions on each ligand is reflected by higher relaxivity values for Gd–PAP–liposomes over the entire frequency range investigated. In fact, both τ_RL and the order parameter, S², assume larger values in this case (Table 1). Furthermore, the faster exchange of the bound water molecules, typical of DTPA-monoamide derivatives, is an additional contributing factor to the increase in r_1p of about 100% at 40 MHz.

It is also possible that, following extrusion and equilibration, the larger Gd–PAP ligand may be preferentially found on the outer leaflet due to steric considerations, so enhancing r₁, relative to Gd–DTPA equivalents. This pre-supposes that water permeation across the membrane is limiting. In the case of limited diffusion of water molecules through the membrane, the relaxivity increases with temperature since the contribution to r₁ of the paramagnetic centers localized in the inner leaflet also increases. In contrast, the relaxivity of both liposomal systems decreases with increasing temperature (see ESI†) and this suggests that their different distribution in the outer-and inner leaflets, if indeed it does occur, is not the main factor that explains their different relaxivity.

Antibody loading

The number of antibody units per liposome was estimated under the assumptions that the size of an immunoliposome was equal to the average hydrodynamic size measured by PCS; here we take an average value of 135 nm, and that the vesicles were unilamellar and spherical.

Firstly, the number of lipid molecules in 1 mL of the final suspensions can be calculated knowing all volumes and assuming quantitative conversion of the original formulation of 62.25 mol% of egg PC (5.2 μmol), 30 mol% of cholesterol (2.5 μmol), 5 mol% of DSPE–PEG–methoxy (0.416 μmol), 1.75 mol% Gd–DTPA–BSA (0.156 μmol) and 0.5 mol% of FITC–PE (0.0416 μmol) to be ca. 5.3 × 10¹³ molecules per mL.³³ The average lipid head group area (weighted by the mole fractions), was calculated to be 0.50 nm², see Table 2. This gives an estimate of 212 578 lipids for each 135 nm liposome. Using this footprint and the liposome concentration we can also estimate that there are 1.52 × 10¹⁶ liposomes per mL.

Table 2 Dimension of lipid head groups³⁵

Lipid	Component	Area [nm^2]
Egg PC	Head group	0.55
DSPE	Head group	0.55
PE	Head group	0.55
Cholesterol	Head group	0.40
DTPA	Head group	0.70

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The antibody (150 kDa) concentration in a solution of purified immunoliposomes was determined using the micro Lowry assay with Peterson's modification to be 103.18 μg mL⁻¹, or 6.9 × 10⁻¹⁰ moles per mL of anti-CD138-conjugated-liposome solution, or ∼4.14 × 10¹⁷ antibody molecules. Hence we can estimate that there are on average ∼27 antibody molecules per liposome (=4.94 × 10¹⁷ antibody molecules/1.52 × 10¹⁶ liposomes).

The estimates of the numbers of molecules and liposomes are subject to the assumption of quantitative conversion of all the components in the formulation. However, that is not the case for the average number of ∼27 antibody molecules per liposome: we have determined the gadolinium incorporation to be 60%, by ICP-AES. The micro Lowry assay on the liposomes demonstrated 52% yield of antibody. Gravimetric analysis, for a range of lipid formulations prepared in our laboratory using the methods described, consistently gives yields of 60–70%.³⁴ This demonstrates that the fractions of all the components incorporated are unchanged from the starting formulation, and the yield from the contrast immunoliposomes preparation process is in the range of 60% (mostly associated with extrusion). Hence it is reasonable to conclude that the number of antibody molecules per liposome is not significantly altered by the reduced yield after processing.

Finally, it was also determined that size of contrast liposomes increased after conjugation with anti-CD138 antibody. It was found that hydrodynamic diameter (from PCS) shifted from 158.6 nm to 161.7 nm for anti-CD138–Gd–DTPA–BSA–liposomes and from 193.9 nm to 208.8 nm for anti-CD138–Gd–PAP–liposomes. Assuming an antibody size of 18 nm, one would expect addition of ca. 27 antibodies to produce a marginal increase in hydrodynamic size. PCS is sensitive to the extent of hydration and is very sensitive to the larger liposomes in the size distribution, so it is not possible to confirm the number of conjugated antibodies by this technique. However, the observations are consistent with significant conjugation, as is shown by the functional study below.

Phantoms of contrast Gd–DTPA–BSA– and Gd–PAP–liposomes were then assessed by T₁-weighted MRI at 1.5 T (Fig. 3). Significant intensity enhancement is observed for Gd–PAP–liposomes as compared to Gd–DTPA–BSA–liposomes, at the same Gd–lipid concentration of 1.75 mol%. The signal of the Gd–DTPA–BSA–liposomes was similar to that obtained for pure water. This demonstrates the possibility of generating contrast under T₁-weighting conditions in MRI. Phantoms of anti-CD138–Gd–DTPA–BSA- and anti-CD138–Gd–PAP-conjugated-liposomes were also imaged with T₁-weighting. Again, it is apparent from Fig. 4, that the signal of anti-CD138–Gd–PAP-conjugated-liposomes is significantly enhanced relative to anti-CD138–Gd–DTPA–BSA-conjugated-liposomes, in this case at the same Gd concentration of 1.75 mol%. The signal generated by the anti-CD138–Gd–DTPA–BSA-conjugated-liposomes was similar to that given by the control (pure water). This MRI analysis demonstrates that significant contrast can be generated using both contrast liposomes and contrast immunoliposomes under imaging conditions and, hence, the presence of the antibody at an average of 27 units per liposome does not compromise the key physicochemical parameters determining the contrast efficiency of the targeted carriers.

Fig. 3 T₁-weighted MRI images of contrast liposomes containing 1.75 mol% (w/v) Gd–liposomes. Gd–PAP–liposomes (sample 1), pure water (sample 2) and Gd–DTPA–BSA–liposomes (sample 3). The imaging field was 1.5 T.

Fig. 4 T₁-weighted MRI images of pure water (sample 1), contrast anti-CD138-conjugated-liposomes, containing 1.75% (w/v) of Gd–DTPA–BSA (sample 2), and of Gd–PAP (sample 3). The imaging field was 1.5 T.

Immuno-liposome performance

Subsequently, human multiple myeloma cancer cell lines (U266 and RPMI8226) were tested for CD138 protein expression just before the performance of in vitro studies using anti-CD138-contrast liposomes. It was estimated that U266 human multiple myeloma cancer cells express CD138 at a significantly (p < 0.05) higher level (approximately 26 fold) in comparison to RPMI8226 and L428 Hodgkin's disease-derived cell lines (Fig. 5). The Hodgkin's disease-derived L428 cell line does not express CD138. Human Burkitt's lymphoma (RAMOS) cells and L428 were tested for CD20 antigen expression with the Rituximab antibody. However, no binding was observed using 1 μg mL⁻¹ of Rituximab antibody with the L428 cell line (Fig. 6). The RAMOS cell line was positively stained using these three different Rituximab concentrations. Thus, it was shown that the RAMOS cell line is CD20-positive and L428 is CD20-negative.

Fig. 5 Examination of antigen expression and antibody binding. U266 multiple myeloma cells demonstrated much higher staining with anti-CD138 antibodies (diagram A) than RPMI8226 (diagram B) or L428 cell lines compared to isotype control antibody (diagram C). It was estimated, using Student's t-test, that there was a significant difference (p value < 0.05) between U266 cells incubated with isotype control antibody and U266 cells incubated with FITC-labelled anti-CD138 antibody (diagram D). Analyses were performed in triplicate and the results were normalized for comparison purposes with plain cells and isotype control-antibody (bottom diagram). The results are the mean ± the standard deviation and the experiment was performed in triplicate. *** Significantly different (Student's t-test, p < 0.001) from control (FITC-labelled mouse IgG κ isotype antibody).

Fig. 6 FACS analyses of RAMOS (human Burkitt's lymphoma) and L428 (Hodgkin's disease-derived) cell lines for CD20 protein expression. RAMOS and L428 cell lines were treated with 1 μg mL⁻¹ of Rituximab (Rit) or Avastin® as a negative control antibody. A 1 : 2000 dilution of Cy5-conjugated anti-human IgG was used as the secondary antibody. RAMOS and L428 cells were also treated with secondary antibody alone as a second negative control.

Studies on in vitro interactions of contrast anti-CD138-conjugated liposomes with cancer cell lines were also performed and visualized by flow cytometry. Approximately 5 × 10⁵ cells were incubated with 0.1 mg mL⁻¹ of liposomes in ‘serum-free’ medium for 1–1.5 hours at 37 °C. Gd–PAP (1.75 mol%) was used in contrast anti-CD138-conjugated liposome preparation. The same amount of contrast lipid was used to generate unconjugated liposomes (liposomes without antibody attached) and as a negative control Herceptin®-conjugated-contrast liposomes. All preparations were washed three times with 0.1 M PBS, pH 7.3.

It was found that anti-CD138–Gd–PAP-conjugated-liposomes bound approximately 2.1 fold more strongly to U266 human multiple myeloma cells than did control liposomes (Herceptin®-conjugated-Gd–PAP–liposomes) (Fig. 7). This suggests that fluorescently-labelled contrast liposomes bind more significantly to U266 human multiple myeloma cell line compared to do negative control liposomes.

Fig. 7 In vitro studies of anti-CD138–Gd–PAP-conjugated liposomes. Anti-CD138–Gd–PAP-conjugated-liposomes bound significantly more to U266 of human multiple myeloma cell lines compared to the negative control (Herceptin®–liposomes; liposomes with isotype control antibody attached) and to plain liposomes (liposomes without antibody attached). Analyses were performed in triplicate and the results were normalized for comparison purposes with blank liposomes and Herceptin®-antibody-conjugated-liposomes. The results are the mean ± the standard deviation. * Significantly different from control (Herceptin®–liposomes; Student's t-test, 0.01 < p < 0.05).

In vitro interactions of fluorescently-labelled Rituximab-conjugated-liposomes with human Burkitt's lymphoma were analysed using flow cytometry (Fig. 8). The RAMOS cell line bound Rituximab-conjugated-labelled-liposomes 2.5 times more strongly than the L428 cell line. The Student's t-test confirms a statistically significant difference (0.001 < p < 0.01) in fluorescence between Avastin®-conjugated-liposomes (negative control) and Rituximab-conjugated-liposomes bound to the RAMOS cell line. These findings demonstrate that Rituximab-labelled liposomes bind specifically to the RAMOS cell line.

Fig. 8 In vitro studies demonstrating specific binding of antibody-conjugated-liposomes to cells expressing the associated target antigen. Fluorescent liposomes conjugated with either Rituximab, Avastin® (negative control antibody) or no antibody (plain or non-antibody conjugated liposomes) were incubated with RAMOS and L428 cell lines. Analyses were performed in triplicate and the results were normalized for comparison purposes with non-antibody-conjugated liposomes and Avastin®-antibody-conjugated-liposomes. The results are the mean ± the standard deviation and the experiments were performed in triplicate. ** Significantly different (Student's t-test, 0.01 < p < 0.001) from control (Avastin®-conjugated-liposomes).

Conclusions

We have shown that Gd–PAP–liposomes exhibit significantly increased spin–lattice relaxivity (per gadolinium ion) as compared to equivalent Gd–DTPA–BSA–liposomes. Given the increased number of ions per lipid, Gd–PAP–liposomes may provide strong contrast under T₁-weighting conditions in MRI. The NMRD analysis identified the key physicochemical differences between the two vehicles that give rise to the enhanced relaxivity for Gd–PAP–liposomes. We further demonstrated that the T₁-weighting capability is maintained on conjugation of the monoclonal anti-CD138 antibody to produce contrast immunoliposomes. It was found that anti-CD138–Gd–PAP-conjugated-liposomes bound approximately 2.1 fold strongly to U266 human multiple myeloma cells than did negative control liposomes of Herceptin®–Gd–PAP–conjugated-liposomes. It was also found that fluorescently labelled Rituximab-conjugated-liposomes bound approximately 2.5 fold more strongly to RAMOS human Burkitt's lymphoma than to the negative control. A preliminary confirmation of the possibility of labelling human multiple myeloma cells with the targeted agent was also undertaken. These findings should contribute towards the preparation of novel, minimally invasive contrast agents for diagnosis of human multiple myeloma and human Burkitt's lymphoma.

Acknowledgements

This study was supported by the Irish Cancer Society, Dublin City University, the Mater Misericordiae University Hospital (BKS06EUS) and the EU PHARMATLANTIC grant. Part of this work was carried out in the Pharmaceutical Biotechnology and Nanomedicine Centre at Northeastern University, Boston, USA. This project has received funding from the European Union’s Seventh Framework Programme for research, technological development and demonstration under grant agreement no. 316973.

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Footnote

† Electronic supplementary information (ESI) available. See DOI: 10.1039/c3ra45400b

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