Nanomedicine: Nanotechnology, Biology, and Medicine xx (2014) xxx – xxx nanomedjournal.com
Quick adjustment of imaging tracer payload, for in vivo applications of theranostic nanostructures in the brain Jesús Agulla, PhD a, b, 1 , David Brea, PhD a, c, 1 , Bárbara Argibay, MsC a , Mercedes Novo d , Francisco Campos a , Tomás Sobrino a , Miguel Blanco a , José Castillo a , Pedro Ramos-Cabrer a,⁎ a
Department of Neurology, Neurovascular Area, Clinical Neurosciences Research Laboratory, Hospital Clínico Universitario, Health Research Institute of Santiago de Compostela (IDIS), University of Santiago de Compostela, Santiago de Compostela, Spain b Research Unit, University Hospital of Salamanca and Institute of Health Sciences of Castilla and Leon, Salamanca, Spain c Cellular and Molecular Neurobiology Research Group and Grup de Recerça en Neurociencies del IGTP, Department of Neurosciences, Fundació Institut d'Investigació en Ciències de la Salut Germans Trias I Pujol-Universitat Autónoma de Barcelona, Badalona, Spain d Single Molecule Fluorescence Research Unit, Department of Physical Chemistry, Faculty of Science, University of Santiago de Compostela, Lugo, Spain Received 8 July 2013; accepted 13 December 2013
Abstract In order to provide sufficient sensibility for detection, selection of an adequate payload of imaging probe is critical, during the design of MRI theranostic nanoplatforms. This fact is particularly crucial for in vivo applications in the brain, where delivery of macromolecules is limited by the blood–brain barrier. Here we report a simple and quick process for the estimation of adequate payloads of gadolinium in liposomes with potential to act as theranostic agents, for in vivo MRI applications in the brain. Our studies show that an excessive payload of gadolinium in liposomes may actually have a negative influence on in vivo T1 contrast. By preparing and characterizing 4 different liposomal compositions of increasing Gadolinium loads, we show that a superior sensitivity for in vivo detection of MRI theranostic molecules can be quickly improved by adjusting the payload of imaging probe in the molecules. © 2013 Elsevier Inc. All rights reserved. Key words: Theranostics; Gadolinium chelates; Probe payload; Contrast quenching; Molecular imaging
During the last decade, application of nanotechnology to the medical field has boosted the development of new and more effective diagnostic and therapeutic tools for the treatment of disease. In particular, the development of molecular nano-
platforms for molecular imaging has enabled the use of conventional imaging techniques to follow up biological processes at molecular level, to reveal new aspects of the mechanisms underlying different pathological processes.
Abbreviations: BBB, Blood–brain Barrier; DLS, Dynamic Light Scattering; DSPC, 1,2-dioctadecanoyl-sn-glycero-3-phosphocholine; FOV, Field of view; Gd, Gadolinium; Gd-DTPA, Gadolinium diethylenetriaminepenta-acetate; Gd-DTPA-BSA, diethylenetriaminepentaacetic acid-bis(stearylamide) gadolinium salt; HBS, Hepes Buffered Solution; MRI, Magnetic Resonance Imaging; NaCl, Sodium Cholride; PEG, Poly-ethylene glycol; r1, Longitudinal relaxivity; r2, Transversal relaxivity; R1, Longitudinal relaxation rate; R2, Transversal relaxation rate; T1, Longitudinal relaxation time; T2, Transversal relaxation time; TE, Echo time; TR, Repetition time; TEM, Transmission electronic microscopy; (v/v), volume/volume ratio; (w/w), mass/mass ratio. Conflict of interests: There are no competing interests present in reference to this work. Financial support: This work has been supported by grants of the Instituto de Salud Carlos III of the Spanish Ministry of Health (projects PI11/02161 and CP09/00074), the Spanish Ministry of Economy and Competence (projects SAF2008-02190, SAF2011-30517 and RETICS-INVICTUS R012/0014) and the European Union (FEDER program). Research contracts from the Miguel Servet program (PRC and TS) and the Sara Borrell program (JA and DB) of the Instituto de Salud Carlos III are deeply acknowledged. ⁎Corresponding author at: Laboratorio de Neurociencias Clínicas. Hospital Clínico Universitario. Travesa da Choupana s/n. 15706 Santiago de Compostela. Spain. E-mail address: [email protected] (P. Ramos-Cabrer). 1 These two authors have contributed equally to this work. 1549-9634/$ – see front matter © 2013 Elsevier Inc. All rights reserved. http://dx.doi.org/10.1016/j.nano.2013.12.004 Please cite this article as: Agulla J., et al., Quick adjustment of imaging tracer payload, for in vivo applications of theranostic nanostructures in the brain.... Nanomedicine: NBM 2014;xx:1-8, http://dx.doi.org/10.1016/j.nano.2013.12.004
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Recently, the concepts of molecular imaging and controlled drug release fields have been merged to rise a new and exciting discipline known as theranostics. 1 In nanotechnology, macromolecular structures with molecular recognition capabilities can be designed to enable the delivery of therapeutics at targeted sites. The inclusion of imaging probes in the same nano-platforms does also allow the real time monitoring of these processes, enabling a rapid adaptation of therapies to the magnitude of the changes observed (i.e. personalized medicine). Some impressive examples of theranostic systems have been reported for cancer, 2 atherosclerosis, 3 and gene therapy, 4 among others. With the progressive aging of population in developed countries, the incidence of neurological diseases is rapidly increasing, 5 and theranostics may represent a decisive tool for the development of novel treatments in this field. However, the use of theranostic approaches for targets located inside the brain is so far rare, due to the difficulties experienced by systemically administered substances to cross the blood–brain barrier (BBB), the functional structure that protects the brain from foreign entities. This is an unresolved issue that requires further study before the concept of theranostics could be widely used in neurosciences. This fact has to be taken into account for the design of macromolecules for in vivo applications in the brain. Though numerous types of nanosystems (e.g. peptides, dendrimers or liposomes) have potential to be used as theranostic agents, 6–13 not all of them are able to cross the BBB in sufficient amounts to offer appropriate sensitivity for in vivo detection. In this sense liposomes are among the most popular nano-platforms for drug delivery to the brain, as they are well-known biocompatible drug carriers, with capacity to effectively cross the blood–brain barrier. 14,15 The inclusion of lanthanides (mostly gadolinium chelates) in the structure of liposomes enables their detection by Magnetic Resonance Imaging (MRI), one of the most used in vivo imaging modalities, due to its non-invasiveness, contrast versatility and high spatial and temporal resolution. However, sensitivity of MRI is low, in comparison to other molecular imaging techniques such as fluorescence and optical imaging, or positron emission tomography, and the use of important amounts of lanthanide nuclei per liposome unit is usually required, in order to achieve high sensibility with these imaging probes. 16,17 The use of large amounts of gadolinium nuclei may rise some concerns about toxicity issues, a fact that will be discussed later. Many are the factors, at molecular level, that affect the contrast achieved by super-paramagnetic substances in vivo, such as the electronic relaxation of the molecule at a given magnetic field, probe concentration, hydration numbers, rotational diffusion, water exchange ratios, pH, temperature, association to proteins, etc. 18 Thus, the (somehow) reasonable assumption that “the higher the payload of contrast agent included in a nano-platform, the higher the signal induced by it”, may actually be adventured for gadolinium-based theranostic agents, where a tight packaging of gadolinium units in a macromolecule of limited size may result in partial quenching of the paramagnetic effect, yielding a poor in vivo contrast, a fact that has been already depicted by other authors. 19
The limited access of molecules to the brain parenchyma through the blood brain barrier, together with the reduced sensitivity of MRI, compels the issue of the selection of a proper payload of contrast agent in molecular imaging nanoplatforms in a key aspect, for the design of these molecules. In this work we report a simple and quick method for the selection of a proper payload of gadolinium chelates in MRIbased macromolecules with potential to act as theranostic agents, showing that it is possible to achieve superior in vivo contrast when the load of gadolinium chelates is properly adjusted.
Methods Preparation of liposomes Liposomes were prepared by the known method of lipid film rehydration. 20,21 Using a 6:1 (v/v) chloroform/methanol mixture (Panreac Química, Barcelona, Spain) as solvent, solutions of phospholipids (typical total amount of lipids of 24 μmol) were prepared including cholesterol (molar fraction of x = 0.35), 1,2-dioctadecanoyl-sn-glycero-3-phosphocholine (DSPC), at four different molar fractions; x = 0.36, x = 0.48, x = 0.54 or x = 0.57, and diethylenetriaminepentaacetic acidbis(stearylamide) gadolinium salt (Gd-DTPA-BSA), in four different molar fractions x = 0.24, x = 0.12, x = 0.06 or x = 0.03, respectively. Notice that the total molar fraction of these two phospholipids was always maintained constant at x = 0.6, but liposomes contained relative 8x, 4x, 2x and 1x amounts of gadolinium lipids. Finally, a molar fraction of x = 0.05 of 1,2distearoyl-sn-glycero-3-phosphoethanolamine-N-[amino (poly ethylene glycol)-2000] ammonium salt (PEG-DSPE) was also added, together with a 0.1% (w/w) of the fluorescent phospholipid 1,2-dioleoyl-sn-glycero-3-phosphoethanolamineN-(lissamine rhodamine B sulfonyl) ammonium salt (Liss Rhodamine PE). All products were obtained from Avanti Polar Lipids, Alabaster, AL, USA. After rotary evaporation of the organic solvent at 30 °C, and overnight drying under a nitrogen current, formed lipid films were hydrated in HEPES-buffered saline (HBS, 10 mM HEPES, 135 mM NaCl, pH 7.4) at 65 °C. Liposomes where then extruded in a Lipex Extruder (Northern Lipids Inc, Burnaby, BC, Canada), keeping the temperature continuously above T = 65 °C (T N Tc for these liposomes) at any time. Liposomes were respectively extruded 2 + 4 + 8 times though 400, 200 and 100 nm polycarbonate membranes (Whatman plc, Kent, UK). Characterization of liposomes: Total lipid concentrations were determined for the 4 different liposome formulations by quantification of phosphates according to Rouser et al. 22 Distributions of liposome sizes in solution were determined at 23 °C by dynamic light scattering (DLS, ZetaSizer NanoS, Malvern Instruments, Worcestershire, UK). For cryo transmission electronic microscopy (cryo-TEM) selected samples were vitrified on carbon-coated using a Vitrobot Mark III (FEI, OR, USA), and 25000X images were acquired on a Tecnai 20 Sphera TEM instrument (FEI, OR, USA) equipped with a LaB6 filament (200 kV) and Gatan cryoholder (− 170 °C).
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Figure 1. Schematic representation of the four different sets of liposomes prepared. All four solutions contain the same total amount of Gadolinium atoms, but the relative number of Gd units per liposome is 1x, 2x, 4x and 8x for the different liposomal formulation.
Animal studies
MR imaging
All experimental protocols involving the use of research animals have been approved by our institutional ethics committee, and were performed according to the guidelines of the Animal Welfare Committee of the host institution and in accordance with applicable legislation of the European Union (DC 86/609/EEC, 2003/65/EC, 2010/637EU). A total of 6 male Sprague–Dawley rats (Harlan Laboratories, Barcelona, Spain) weighing 352 ± 36 g were used in this work. Animals were kept in a controlled environment at 22 ± 1 °C and 60 ± 5% humidity, with 12/12 h light/darkness cycles. Animals were fed ad libitum with standard diet pellets and tap water. All surgical procedures and MRI studies were conducted under Sevofluorane (Abbott laboratories, IL, USA) anesthesia at a level of 3-4% using a carrier 65:35 gas mixture of N20:02. To study the in vivo contrast introduced by the designed molecules in the rat brain we performed intraparenchymal injections of 10 μl of different solutions of liposomes in PBS, through a hole buried in the skull of the animals. In this study we intend to show the potential of the agent to induce contrast, depending on its composition, therefore we decided to “bypass” the problematic of crossing the Blood–brain Barrier, for the sake of simplicity. The problem of achieving sufficient concentration in the brain parenchyma after systemic injections (trespassing the BBB) will be the subject of future studies.
Magnetic resonance imaging studies were conducted on a 9.4 T MR system (Bruker Biospin, Ettlingen, Germany) with 440 mT/m gradients, and using a combination of a linear birdcage resonator of 7 cm of diameter, for signal transmission, and a 2x2 arrayed surface coil for signal detection (both from Bruker Biospin, Ettlinglen, Germany). T1 weighted images (T1w) were typically acquired using a RARE (fast spin echo) sequence (rare factor = 4, averages = 2) with an effective echo time of TE = 22 ms and repetition times of TR = 900 ms (flip angle = 90°). Whole brain was covered by 14 consecutive axial slices of 1 mm thickness, with a field-ofview (FOV) of 19.2x19.2 mm (with saturation bands to suppress signal outside this FOV) and an in-plane resolution of 100 microns (192x192 points matrix). Coronal images where acquired using a FOV of 32x19.2 mm, keeping the same spatial resolution. FOV was adapted to required sizes for MR imaging of solutions (placed on 0.2 ml Eppendorf tubes) on in vitro studies. For the obtaining of T1 maps the same RARE sequence was used to follow a saturation-recovery approach, 23 acquiring series of 12 images with repetition times ranging from 250 ms to 12 s (following an exponential distribution). Proton density maps (with certain T2 weight, at the injection site, introduced by the contrast agent), T1 and T2 relaxation time maps, and R1 and R2 relaxation rate maps were calculated by a
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Figure 2. (A) CryoTEM micrograph of liposomes showing the formation of (mostly) unilamellar small vesicles (SUVs) of uniform size. An amplification of one of the nanostructures allows the visualization of the phospholipid bylayer. (B) Histogram showing the distribution of sizes on a solution of liposomes (hydrodynamic diameters), as measured by Dynamic Light Scattering (DLS). A narrow distribution around 130 nm was determined at 23 °C.
pixel-by-pixel mono-exponential fitting of the data, with selfdeveloped applications for Image-J (Rasband, W.S., ImageJ, U. S. National Institutes of Health, Bethesda, Maryland, USA, http://imagej.nih.gov/ij/, 1997–2012), or using ParaVision 5.1 (Bruker Biospin, Ettlingen, Germany). Longitudinal relaxivities (r1) of the different liposomal compositions were obtained from the slope of linear plots obtained by representing the corresponding relaxation rates (R1 = 1/T1) of 12 solutions of different concentration of liposomes versus the total concentration of gadolinium ([Gd]total) on those solutions. For the construction of these plots, sets of 12 seriated dilutions of the liposomal compositions were prepared by the dilution-extraction method, 24 using stock solutions of liposomes containing (typically) 5–40 μmol lipids/ml, i.e. [Gd]total = 90 μM, with a dilution factor of Vf = 0.8, achieved by using a stock solution of V = 1 ml and adding and extracting samples of v = 0.25 ml, stored on Eppendorf tubes that were directly introduced on the MR system. Sample preparation For the characterization of the MR properties of magnetoliposomes, we first prepared stock solutions for each of the 4 synthesized liposomal compositions. Since (gadolinium-phospholipids/total amount of phospholipids) ratios were different for the 4 liposomal compositions (x = 0.24, x = 0.12, x = 0.06 and
x = 0.03, for gadolinium-doped phospholipids, giving relative gadolinium amounts of 8x, 4x, 2x and 1x), stock solutions were prepared with respective 1x, 2x, 4x and 8x concentrations of liposomes (i.e. 5, 10, 20 and 40 μmol lipids/ml), in order to achieve the same absolute amount of gadolinium in solution ([Gd]total = 90 μM), as schematically represented in Figure 1. Then, each stock solution was used to prepare 12 seriated dilutions, by using the dilution-extraction method. 24 This method consists on the dilution of the stock solution with a small volume of solvent, retiring later exactly the same volume from the diluted solution, keeping the total volume of the stock solution constant after each dilution-extraction step, and using the retired volume of solution as sample for measurement. Thus, after a series of successive dilution-extraction steps, a precise series of sample solutions of decreasing concentrations is obtained, using low volumes of stock solution (which is quite important when using expensive products). In our case we used a volume of V = 1 ml of stock solution for each liposomal composition to prepare the complete series of solutions (12 solutions for each liposomal composition) to be measured. The retired volume on each dilution step was v = 0.25 ml (Vf factor of 0.8), 24 stored on Eppendorf tubes for MRI studies. A final in vivo study was conducted to assess the sensitivity of detection of the prepared liposomal compositions in the brain. For this study 4 rats were injected with v = 10 μl of liposome solutions with liposomal concentrations of 6.25, 12.5, 25 and
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Figure 3. Color-coded T1 and T2 maps of 12 solutions with variable concentration of liposomes (gadolinium content on solution ranging 10–120 μM) obtained at 9.4 Tesla for the composition of liposomes with the highest payload of gadolinium per liposome unit (white numbers in the background represent calculated concentration of gadolinium in each tube in mM). Similar maps where obtained for the other 3 compositions of liposomes.
50 μM, respectively, for liposomal compositions containing 8x, 4X, 2x and 1x relative amounts of gadolinium lipids. Thus, the total amount of gadolinium injected on each animal was the same, Gdtotal = 15 nanomol (v = 10 μl, [Gd]total = 1.5 mM) . An additional rat was injected with v = 10 μl of a solution 50 mM of Gd-DTPA (Dotarem, Geubert, France).
Results The structure of synthesized liposomes was inspected by acquiring cryo-TEM images of frozen solutions of liposomes (Figure 2, A). These pictures show the formation of (mainly) small unilamellar vesicles (SUVs) with a narrow distribution of sizes in solution of 130 nm, as measured by dynamic lightscattering. A typical histogram of hydrodynamic diameters obtained for liposomal formulations (centered at 129.8 nm) is presented in Figure 2, B (similar results were obtained for other compositions tested). Distributions of sizes present a half-width at half-height of circa 25 nm (24.2 nm in Figure 2, B). Liposomes remain stable in solution for a week, when stored at 4 °C (less than 7% of larger populations detected by DLS at day 7 of storage, data not shown), which is a relevant logistic issue when seriated in vivo studies are planned. For the obtaining of T1 relaxation times, samples were prepared by the dilution-extraction method, as described above, and placed on the MR system on groups of n = 6 samples. An example of T1 and T2 parametric maps obtained for a complete series of solutions of one of the liposomal compositions (x = 0.24, relative 8x load of gadolinium) is presented in Figure 3. On this color-coded image, each circle represents a transversal cut through the sample tube, along with the calculated total concentration of gadolinium. Similar results were obtained for the other three liposomal compositions. Finally, relaxation rates of each solution (R1 = 1/T1) were calculated, and plotted against the total concentration of gadolinium for each solutions, obtaining the linear plots presented in Figure 4. The slope of
each of these linear plots represents the longitudinal relaxivity (r1) of the corresponding liposomal composition. The longitudinal relaxivities obtained were r1 = 2.3, 3.4, 3.5 and 4.9 s - 1 mM - 1 for liposomes containing relative 8x, 4x, 2x and 1x molar fractions of Gadolinium-lipids, which are in the order of most T1 commercial contrast agents for clinical use. 18 We have found a value of r1 = 3.60 ± 0.02 mM - 1 s - 1 at 9.4 T (Figure 4) for commercial Gd-DTPA (DOTAREM), which is in agreement with previously published values for this agent at this magnetic field. 25 Interestingly, the liposomal composition with the lowest molar fraction of Gadolinium-lipids on our study (x = 0.03, relative 1X) presented the highest value for the longitudinal relaxivity. For the in vivo study, Proton density images (with certain T2 weight, at the injection site, introduced by the contrast agent) and R1 maps where obtained for each animal (Figure 5) showing that while the injection of the liposomal composition with the highest payload of gadolinium (molar fraction of x = 0.24) is barely distinguishable form the background signal of brain tissue, the composition including the lowest molar fraction of the tested ones (x = 0.03) is clearly visible, and the contrast induced is comparable to that obtained by a 30 times higher dose of commercial Gd-DTPA. It is noteworthy that solutions of liposomes remain closely contained around the injection site (the mean measured volume of brain parenchyma with reduced T1 time for all animals was 12.4 ± 3.9 μl, after injection of 10 μl of contrast agent), conversely to what happens with the injection of free Gd-DTPA, that diffuses through the brain parenchyma (Vreduced T1 = 71.6 μl, Figure 5). This fact may indicate that liposomes do not behave as a simple solutions (as free Gd-DTPA does) and diffusion along the tissue is restrained, to some extent.
Discussion As we have already mentioned, stealth cholesterol/DSPC liposomes were used as structural basis of our theranostic agents.
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Figure 4. Plot of measured longitudinal relaxation rates (R1 = 1/T1) obtained from T1 maps vs. the total concentration of gadolinium on those solutions ([Gd]Total), for Gd-DTPA and for all studied liposomal compositions. The slope of each of these plots corresponds to the relaxivities (r1) of Gd-DTPA and of each liposomes composition set.
We included a fraction of PEG-conjugated lipids to avoid in vivo aggregation, opsonization and rapid clearance from blood by the reticuloendothelial system. 6,26–29 The development of Stealth technology, by conjugation of polyethylene glycol chains to the surface of liposomes, has been proven to extend the circulatory half-life of the molecules for up to 18 h after systemic injection, increasing their biodisponibility in vivo. 19 Indeed, the integrity and biodisponibility of liposomes, once injected in a living bean, is an important issue to consider for potential in vivo applications of liposomes. Liposomes have been described as one of the most stable systems for drug delivery, 30 and it is known that their integrity in vivo depends on their composition and size. Cholesterol-stabilized liposomes with sizes on the range from 70 to 200 nm (like those prepared here) have been shown to remain in the body for periods of time up to 24–48 h after injection. 30 Some authors have even reported residence times as prolonged as 72 h, assuming that liposomes stored in the spleen of the animals, and released into the blood stream afterwards. 25 Thus the preparation of liposomes like those described in this study, with a mean size of 100 nm, PEG grafts on their surface, and stabilized by cholesterol (our DSPC-Cholesterol liposomes preset a critical temperature over 65 °C) should ensure a sufficient residence of these systems in vivo, to achieve a theranostic activity, as foreseen for these molecules. In addition, the limitation of liposomes’ size to circa 100 nm is expected to facilitate their crossing through the blood brain-barrier, 15,31 which becomes relevant when the molecular event to be followed takes place inside the brain parenchyma, such is the case for theranostic strategies on neurological disorders. Following a multimodal approach, 16,17 phospholipids containing rhodamine or gadolinium were both included into liposomes, to make them traceable by fluorescence and MRI techniques, respectively, thus enabling their use as diagnostic tools. The total load of gadolinium per liposome unit is a critical parameter that has to be optimized for theranostic (molecular
imaging) purposes, as it is crucial for the induction of T1 contrast on MR images. Here, we show that the intensity of the T1 effect induced by a particular load of gadolinium in liposomes is not linearly dependent on gadolinium content. Thus, a series of compositions need to be prepared and tested in order to find the most appropriate payload of the superparamagnetic agent, for optimized sensitivity on in vivo applications. In this work we describe a quick and precise experimental procedure to achieve this goal. Our results confirm the fact, already pointed by other authors, 19 that gadolinium paramagnetic activity is not an additive property when multiple units of this lanthanide are located in the structure of liposomes, in respect to a situation on which gadolinium atoms are free in solution. The relaxivity of magneto-liposomes does not continuously rise by a limitless increase of gadolinium units in their structure. 17,18,32 In fact, tight packing of high amounts of gadolinium units in a reduced space may partially quench the paramagnetic effect of gadolinium, yielding lower global relaxivities. 33 It is important to remember that T1 contrast is generated when water molecules have access to the coordination spheres of gadolinium atoms in solutions. Tight packaging of Gd chelates in liposomes hampers the access of water molecules to those coordination spheres. On the other hand, the use of a large fraction of gadoliniumcontaining lipids in liposomes may also affect the rigidity and permeability of the phospholipid bilayer of liposomes to water molecules. 18,34 Permeability of the membrane of liposomes plays an important role in the final relaxivity of macromolecular molecules, as it has been described by Ghaghada et al. 19 On the other hand, relaxivity of macromolecules containing multiple units of gadolinium chelates shows a great dependency on molecular mobility, 18 but in our study liposomes are always constructed in the same way, with same sizes and total amount of phospholipids. Therefore it may be expected that molecular motions would be similar for all of them, and the rigidity and permeability of liposomes to water molecules may be the predominant effect here. Changes on the water residence time inside the inner cavity of liposomes may be also playing a key role in the contras generated by the designed molecules, which causes a more important effect on global contrast than the electronic relaxivity of gadolinium atoms per se. 18,19,33 Last but not least, the number of gadolinium ions per liposome does not just condition the longitudinal relaxivity of the contrast agent (r1, in mM - 1 s - 1) but does also affect to the transversal relaxation rates (R2 = 1/T2) of water molecules, introducing a negative contrast on T2 weighted MR images. Thus, the contrast achieved on T1 weighted images (which do intrinsically have some T2 weight too) 23 is somehow governed by both relaxivities (r1 and r2), rather than by r1 alone. In fact, an excessive payload of gadolinium may be contraindicated, due to a disproportionate weight of T2 effects in the total contrast. 32 The weight of this effect will be highly dependent on the composition of the system. Thus, Courant et al. 35 have described gadolinium-containing hydrogels with r2 values 2.5 time higher than r1 values (at 1.5 T), claiming a powerful dual T1-T2 effect for those systems. Our liposomes presented a mean r2 relaxivity of 8.38 mM - 1 s - 1, at 9.4 T, which represents 2–3 times higher than observed r1 values, therefore T2 weight is expected to play a
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Figure 5. Grey-scaled Proton density MR images (with certain T2 weight, at the injection site, introduced by the contrast agent) and color-coded R1 parametric maps (with magnification of the injection site) of rat brains obtained in vivo, 1 h after intraparenchymal injections of 500 nanomol of Gd-DTPA (one animal) in 10 μl, of four different formulations of magnetoliposomes, containing the same total amount of gadolinium in solution (15 nanomol), but different lipid concentrations (6.25, 12.5, 25 and 50 μM for 8x, 4X, 2x and 1x liposomes composition sets). The lowest payload of gadolinium units per liposomes (liposomes 1x) induced the highest T1 contrast, in vivo, which resulted comparable to Gd-DTPA injections, even using more than 30 times lower amounts of gadolinium in the liposomes respect to the free agent.
similar role on total contrast as that described by Courant et al. 35 In summary, many are the factors that influence the relaxivity of complex macromolecules, and relaxivity must be optimized case by case. 18 A quick method to perform such optimization, like the one presented here, is highly advantageous. The reduction of the total number of gadolinium units by such a high factor, in order to produce a similar contrast in vivo, is very important to minimize toxic effects from gadolinium nuclei. Gadolinium is a very toxic substance in its free state. For this reason, chelation of Gd units by a coligand (such as DTPA) is required for in vivo applications. 18 This process, on the other hand, reduces the number of coordination sites for water (hydration number, q), and therefore the relaxivity of these molecules. Attempts to increase sensitivity of detection by increasing the number of coordination sites will induce weaker links between gadolinium atoms and the chelating agents, increasing the risk of release of free gadolinium into tissues. Therefore, approaches like packing Gd-chelates in liposomes to achieve contrast are preferable, since Gadolinium lipids keep the highly stable chelation of Gd nuclei by diethylenetriamine pentaacetic acid (DTPA) molecules. Additionally, the low molecular weighted contrast agent (GdDTPA) diffuses rapidly through the brain tissue, and may be easily dumped into one of the lateral ventricles of the brain (Figure 5), while liposomes stay concentrated at the injection site. Free diffusion of the contrast agent through the brain tissue is not a desirable effect for a theranostic agent, which is supposed
to remain at the targeted site, in molecular recognition processes. In this sense liposomal based contrast agents offer superior specificity for in vivo applications in the brain. In summary, in this work we have demonstrated that the intensity of the T1 effect induced by a particular load of gadolinium in theranostic (molecular imaging) molecules may be difficult to predict (complex models may be required for that purpose), 25 and just assuming that liposome’s relaxivities would continuously rise by increasing gadolinium units in their structure may have negative consequences, due to the particularities of the physical mechanisms behind T1 contrast induced by lanthanides. The access of water molecules to the coordination spheres of these molecules is crucial, and highly influenced by the number and disposition of gadolinium units on the structure of macromolecules. We have presented a simple, quick and accurate experimental procedure to estimate the relaxivities of MR theranostic nano-platforms based on gadolinium, using minimal amounts of these expensive molecules. Results were confirmed by in vivo studies in the rat brain.
Acknowledgments Authors acknowledge Prof. Dr. Klaas Nicolay and Dr. Gustav Strijkers, From Eindhoven University of Technology (The Netherlands), for sharing their knowledge on the synthesis and characterization of liposomes.
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References 20. 1. Picard FJ, Bergeron MG. Rapid molecular theranostics in infectious diseases. Drug Discov Today 2002;7(21):1092-101. 2. Fernandez-Fernandez A, Manchanda R, McGoron A. Theranostic Applications of Nanomaterials in Cancer: Drug Delivery, Image-Guided Therapy, and Multifunctional Platforms. Appl Biochem Biotechnol 2011;165(7):1628-51. 3. Lobatto ME, Fuster V, Fayad ZA, Mulder WJM. Perspectives and opportunities for nanomedicine in the management of atherosclerosis. Nat Rev Drug Discov 2011;10(11):835-52. 4. Shim MS, Kwon YJ. Stimuli-responsive polymers and nanomaterials for gene delivery and imaging applications. Adv Drug Deliv Rev 2012; 64(11):1046-59. 5. Kinlay S. Changes in Stroke Epidemiology, Prevention, and Treatment. Circulation 2011;124(19):e494-6. 6. Allen TM. Long-circulating (sterically stabilized) liposomes for targeted drug delivery. Trends Pharmacol Sci 1994;15(7):215-20. 7. Drummond DC, Meyer O, Hong K, Kirpotin DB, Papahadjopoulos D. Optimizing liposomes for delivery of chemotherapeutic agents to solid tumors. Pharmacol Rev 1999;51(4):691-743. 8. Kabanov AV. Polymer genomics: an insight into pharmacology and toxicology of nanomedicines. Adv Drug Deliv Rev 2006;58(15):1597-621. 9. Kataoka K, Harada A, Nagasaki Y. Block copolymer micelles for drug delivery: design, characterization and biological significance. Adv Drug Deliv Rev 2001;47(1):113-31. 10. Lee CC, MacKay JA, Frechet JM, Szoka FC. Designing dendrimers for biological applications. Nat Biotechnol 2005;23(12):1517-26. 11. Lu Y, Low PS. Folate-mediated delivery of macromolecular anticancer therapeutic agents. Adv Drug Deliv Rev 2002;54(5):675-93. 12. Nori A, Kopecek J. Intracellular targeting of polymer-bound drugs for cancer chemotherapy. Adv Drug Deliv Rev 2005;57(4):609-36. 13. Widera A, Norouziyan F, Shen WC. Mechanisms of TfR-mediated transcytosis and sorting in epithelial cells and applications toward drug delivery. Adv Drug Deliv Rev 2003;55(11):1439-66. 14. Denora N, Trapani A, Laquintana V, Lopedota A, Trapani G. Recent advances in medicinal chemistry and pharmaceutical technology–strategies for drug delivery to the brain. Curr Top Med Chem 2009;9(2):182-96. 15. Vlieghe P, Khrestchatisky M. Medicinal Chemistry Based Approaches and Nanotechnology-Based Systems to Improve CNS Drug Targeting and Delivery. Med Res Rev 2013;33(3):457-516. 16. Mulder WJ, Griffioen AW, Strijkers GJ, Cormode DP, Nicolay K, Fayad ZA. Magnetic and fluorescent nanoparticles for multimodality imaging. Nanomedicine (Lond) 2007;2(3):307-24. 17. Mulder WJ, Strijkers GJ, van Tilborg GA, Griffioen AW, Nicolay K. Lipid-based nanoparticles for contrast-enhanced MRI and molecular imaging. NMR Biomed 2006;19(1):142-64. 18. Caravan P. Strategies for increasing the sensitivity of gadolinium based MRI contrast agents. Chem Soc Rev 2006;35(6):512-23. 19. Ghaghada K, Hawley C, Kawaji K, Annapragada A, Mukundan Jr S. T1 relaxivity of core-encapsulated gadolinium liposomal contrast agents–
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Quick adjustment of imaging tracer payload, for in vivo applications of theranostic nanostructures in the brain Jesús Agulla, PhD a, b, 1 , David Brea, PhD a, c, 1 , Bárbara Argibay, MsC a , Mercedes Novo d , Francisco Campos a , Tomás Sobrino a , Miguel Blanco a , José Castillo a , Pedro Ramos-Cabrer a,⁎ a
Department of Neurology, Neurovascular Area, Clinical Neurosciences Research Laboratory, Hospital Clínico Universitario, Health Research Institute of Santiago de Compostela (IDIS), University of Santiago de Compostela, Santiago de Compostela, Spain b Research Unit, University Hospital of Salamanca and Institute of Health Sciences of Castilla and Leon, Salamanca, Spain c Cellular and Molecular Neurobiology Research Group and Grup de Recerça en Neurociencies del IGTP, Department of Neurosciences, Fundació Institut d'Investigació en Ciències de la Salut Germans Trias I Pujol-Universitat Autónoma de Barcelona, Badalona, Spain d Single Molecule Fluorescence Research Unit, Department of Physical Chemistry, Faculty of Science, University of Santiago de Compostela, Lugo, Spain Received 8 July 2013; accepted 13 December 2013
Abstract In order to provide sufficient sensibility for detection, selection of an adequate payload of imaging probe is critical, during the design of MRI theranostic nanoplatforms. This fact is particularly crucial for in vivo applications in the brain, where delivery of macromolecules is limited by the blood–brain barrier. Here we report a simple and quick process for the estimation of adequate payloads of gadolinium in liposomes with potential to act as theranostic agents, for in vivo MRI applications in the brain. Our studies show that an excessive payload of gadolinium in liposomes may actually have a negative influence on in vivo T1 contrast. By preparing and characterizing 4 different liposomal compositions of increasing Gadolinium loads, we show that a superior sensitivity for in vivo detection of MRI theranostic molecules can be quickly improved by adjusting the payload of imaging probe in the molecules. © 2013 Elsevier Inc. All rights reserved. Key words: Theranostics; Gadolinium chelates; Probe payload; Contrast quenching; Molecular imaging
During the last decade, application of nanotechnology to the medical field has boosted the development of new and more effective diagnostic and therapeutic tools for the treatment of disease. In particular, the development of molecular nano-
platforms for molecular imaging has enabled the use of conventional imaging techniques to follow up biological processes at molecular level, to reveal new aspects of the mechanisms underlying different pathological processes.
Abbreviations: BBB, Blood–brain Barrier; DLS, Dynamic Light Scattering; DSPC, 1,2-dioctadecanoyl-sn-glycero-3-phosphocholine; FOV, Field of view; Gd, Gadolinium; Gd-DTPA, Gadolinium diethylenetriaminepenta-acetate; Gd-DTPA-BSA, diethylenetriaminepentaacetic acid-bis(stearylamide) gadolinium salt; HBS, Hepes Buffered Solution; MRI, Magnetic Resonance Imaging; NaCl, Sodium Cholride; PEG, Poly-ethylene glycol; r1, Longitudinal relaxivity; r2, Transversal relaxivity; R1, Longitudinal relaxation rate; R2, Transversal relaxation rate; T1, Longitudinal relaxation time; T2, Transversal relaxation time; TE, Echo time; TR, Repetition time; TEM, Transmission electronic microscopy; (v/v), volume/volume ratio; (w/w), mass/mass ratio. Conflict of interests: There are no competing interests present in reference to this work. Financial support: This work has been supported by grants of the Instituto de Salud Carlos III of the Spanish Ministry of Health (projects PI11/02161 and CP09/00074), the Spanish Ministry of Economy and Competence (projects SAF2008-02190, SAF2011-30517 and RETICS-INVICTUS R012/0014) and the European Union (FEDER program). Research contracts from the Miguel Servet program (PRC and TS) and the Sara Borrell program (JA and DB) of the Instituto de Salud Carlos III are deeply acknowledged. ⁎Corresponding author at: Laboratorio de Neurociencias Clínicas. Hospital Clínico Universitario. Travesa da Choupana s/n. 15706 Santiago de Compostela. Spain. E-mail address: [email protected] (P. Ramos-Cabrer). 1 These two authors have contributed equally to this work. 1549-9634/$ – see front matter © 2013 Elsevier Inc. All rights reserved. http://dx.doi.org/10.1016/j.nano.2013.12.004 Please cite this article as: Agulla J., et al., Quick adjustment of imaging tracer payload, for in vivo applications of theranostic nanostructures in the brain.... Nanomedicine: NBM 2014;xx:1-8, http://dx.doi.org/10.1016/j.nano.2013.12.004
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Recently, the concepts of molecular imaging and controlled drug release fields have been merged to rise a new and exciting discipline known as theranostics. 1 In nanotechnology, macromolecular structures with molecular recognition capabilities can be designed to enable the delivery of therapeutics at targeted sites. The inclusion of imaging probes in the same nano-platforms does also allow the real time monitoring of these processes, enabling a rapid adaptation of therapies to the magnitude of the changes observed (i.e. personalized medicine). Some impressive examples of theranostic systems have been reported for cancer, 2 atherosclerosis, 3 and gene therapy, 4 among others. With the progressive aging of population in developed countries, the incidence of neurological diseases is rapidly increasing, 5 and theranostics may represent a decisive tool for the development of novel treatments in this field. However, the use of theranostic approaches for targets located inside the brain is so far rare, due to the difficulties experienced by systemically administered substances to cross the blood–brain barrier (BBB), the functional structure that protects the brain from foreign entities. This is an unresolved issue that requires further study before the concept of theranostics could be widely used in neurosciences. This fact has to be taken into account for the design of macromolecules for in vivo applications in the brain. Though numerous types of nanosystems (e.g. peptides, dendrimers or liposomes) have potential to be used as theranostic agents, 6–13 not all of them are able to cross the BBB in sufficient amounts to offer appropriate sensitivity for in vivo detection. In this sense liposomes are among the most popular nano-platforms for drug delivery to the brain, as they are well-known biocompatible drug carriers, with capacity to effectively cross the blood–brain barrier. 14,15 The inclusion of lanthanides (mostly gadolinium chelates) in the structure of liposomes enables their detection by Magnetic Resonance Imaging (MRI), one of the most used in vivo imaging modalities, due to its non-invasiveness, contrast versatility and high spatial and temporal resolution. However, sensitivity of MRI is low, in comparison to other molecular imaging techniques such as fluorescence and optical imaging, or positron emission tomography, and the use of important amounts of lanthanide nuclei per liposome unit is usually required, in order to achieve high sensibility with these imaging probes. 16,17 The use of large amounts of gadolinium nuclei may rise some concerns about toxicity issues, a fact that will be discussed later. Many are the factors, at molecular level, that affect the contrast achieved by super-paramagnetic substances in vivo, such as the electronic relaxation of the molecule at a given magnetic field, probe concentration, hydration numbers, rotational diffusion, water exchange ratios, pH, temperature, association to proteins, etc. 18 Thus, the (somehow) reasonable assumption that “the higher the payload of contrast agent included in a nano-platform, the higher the signal induced by it”, may actually be adventured for gadolinium-based theranostic agents, where a tight packaging of gadolinium units in a macromolecule of limited size may result in partial quenching of the paramagnetic effect, yielding a poor in vivo contrast, a fact that has been already depicted by other authors. 19
The limited access of molecules to the brain parenchyma through the blood brain barrier, together with the reduced sensitivity of MRI, compels the issue of the selection of a proper payload of contrast agent in molecular imaging nanoplatforms in a key aspect, for the design of these molecules. In this work we report a simple and quick method for the selection of a proper payload of gadolinium chelates in MRIbased macromolecules with potential to act as theranostic agents, showing that it is possible to achieve superior in vivo contrast when the load of gadolinium chelates is properly adjusted.
Methods Preparation of liposomes Liposomes were prepared by the known method of lipid film rehydration. 20,21 Using a 6:1 (v/v) chloroform/methanol mixture (Panreac Química, Barcelona, Spain) as solvent, solutions of phospholipids (typical total amount of lipids of 24 μmol) were prepared including cholesterol (molar fraction of x = 0.35), 1,2-dioctadecanoyl-sn-glycero-3-phosphocholine (DSPC), at four different molar fractions; x = 0.36, x = 0.48, x = 0.54 or x = 0.57, and diethylenetriaminepentaacetic acidbis(stearylamide) gadolinium salt (Gd-DTPA-BSA), in four different molar fractions x = 0.24, x = 0.12, x = 0.06 or x = 0.03, respectively. Notice that the total molar fraction of these two phospholipids was always maintained constant at x = 0.6, but liposomes contained relative 8x, 4x, 2x and 1x amounts of gadolinium lipids. Finally, a molar fraction of x = 0.05 of 1,2distearoyl-sn-glycero-3-phosphoethanolamine-N-[amino (poly ethylene glycol)-2000] ammonium salt (PEG-DSPE) was also added, together with a 0.1% (w/w) of the fluorescent phospholipid 1,2-dioleoyl-sn-glycero-3-phosphoethanolamineN-(lissamine rhodamine B sulfonyl) ammonium salt (Liss Rhodamine PE). All products were obtained from Avanti Polar Lipids, Alabaster, AL, USA. After rotary evaporation of the organic solvent at 30 °C, and overnight drying under a nitrogen current, formed lipid films were hydrated in HEPES-buffered saline (HBS, 10 mM HEPES, 135 mM NaCl, pH 7.4) at 65 °C. Liposomes where then extruded in a Lipex Extruder (Northern Lipids Inc, Burnaby, BC, Canada), keeping the temperature continuously above T = 65 °C (T N Tc for these liposomes) at any time. Liposomes were respectively extruded 2 + 4 + 8 times though 400, 200 and 100 nm polycarbonate membranes (Whatman plc, Kent, UK). Characterization of liposomes: Total lipid concentrations were determined for the 4 different liposome formulations by quantification of phosphates according to Rouser et al. 22 Distributions of liposome sizes in solution were determined at 23 °C by dynamic light scattering (DLS, ZetaSizer NanoS, Malvern Instruments, Worcestershire, UK). For cryo transmission electronic microscopy (cryo-TEM) selected samples were vitrified on carbon-coated using a Vitrobot Mark III (FEI, OR, USA), and 25000X images were acquired on a Tecnai 20 Sphera TEM instrument (FEI, OR, USA) equipped with a LaB6 filament (200 kV) and Gatan cryoholder (− 170 °C).
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Figure 1. Schematic representation of the four different sets of liposomes prepared. All four solutions contain the same total amount of Gadolinium atoms, but the relative number of Gd units per liposome is 1x, 2x, 4x and 8x for the different liposomal formulation.
Animal studies
MR imaging
All experimental protocols involving the use of research animals have been approved by our institutional ethics committee, and were performed according to the guidelines of the Animal Welfare Committee of the host institution and in accordance with applicable legislation of the European Union (DC 86/609/EEC, 2003/65/EC, 2010/637EU). A total of 6 male Sprague–Dawley rats (Harlan Laboratories, Barcelona, Spain) weighing 352 ± 36 g were used in this work. Animals were kept in a controlled environment at 22 ± 1 °C and 60 ± 5% humidity, with 12/12 h light/darkness cycles. Animals were fed ad libitum with standard diet pellets and tap water. All surgical procedures and MRI studies were conducted under Sevofluorane (Abbott laboratories, IL, USA) anesthesia at a level of 3-4% using a carrier 65:35 gas mixture of N20:02. To study the in vivo contrast introduced by the designed molecules in the rat brain we performed intraparenchymal injections of 10 μl of different solutions of liposomes in PBS, through a hole buried in the skull of the animals. In this study we intend to show the potential of the agent to induce contrast, depending on its composition, therefore we decided to “bypass” the problematic of crossing the Blood–brain Barrier, for the sake of simplicity. The problem of achieving sufficient concentration in the brain parenchyma after systemic injections (trespassing the BBB) will be the subject of future studies.
Magnetic resonance imaging studies were conducted on a 9.4 T MR system (Bruker Biospin, Ettlingen, Germany) with 440 mT/m gradients, and using a combination of a linear birdcage resonator of 7 cm of diameter, for signal transmission, and a 2x2 arrayed surface coil for signal detection (both from Bruker Biospin, Ettlinglen, Germany). T1 weighted images (T1w) were typically acquired using a RARE (fast spin echo) sequence (rare factor = 4, averages = 2) with an effective echo time of TE = 22 ms and repetition times of TR = 900 ms (flip angle = 90°). Whole brain was covered by 14 consecutive axial slices of 1 mm thickness, with a field-ofview (FOV) of 19.2x19.2 mm (with saturation bands to suppress signal outside this FOV) and an in-plane resolution of 100 microns (192x192 points matrix). Coronal images where acquired using a FOV of 32x19.2 mm, keeping the same spatial resolution. FOV was adapted to required sizes for MR imaging of solutions (placed on 0.2 ml Eppendorf tubes) on in vitro studies. For the obtaining of T1 maps the same RARE sequence was used to follow a saturation-recovery approach, 23 acquiring series of 12 images with repetition times ranging from 250 ms to 12 s (following an exponential distribution). Proton density maps (with certain T2 weight, at the injection site, introduced by the contrast agent), T1 and T2 relaxation time maps, and R1 and R2 relaxation rate maps were calculated by a
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Figure 2. (A) CryoTEM micrograph of liposomes showing the formation of (mostly) unilamellar small vesicles (SUVs) of uniform size. An amplification of one of the nanostructures allows the visualization of the phospholipid bylayer. (B) Histogram showing the distribution of sizes on a solution of liposomes (hydrodynamic diameters), as measured by Dynamic Light Scattering (DLS). A narrow distribution around 130 nm was determined at 23 °C.
pixel-by-pixel mono-exponential fitting of the data, with selfdeveloped applications for Image-J (Rasband, W.S., ImageJ, U. S. National Institutes of Health, Bethesda, Maryland, USA, http://imagej.nih.gov/ij/, 1997–2012), or using ParaVision 5.1 (Bruker Biospin, Ettlingen, Germany). Longitudinal relaxivities (r1) of the different liposomal compositions were obtained from the slope of linear plots obtained by representing the corresponding relaxation rates (R1 = 1/T1) of 12 solutions of different concentration of liposomes versus the total concentration of gadolinium ([Gd]total) on those solutions. For the construction of these plots, sets of 12 seriated dilutions of the liposomal compositions were prepared by the dilution-extraction method, 24 using stock solutions of liposomes containing (typically) 5–40 μmol lipids/ml, i.e. [Gd]total = 90 μM, with a dilution factor of Vf = 0.8, achieved by using a stock solution of V = 1 ml and adding and extracting samples of v = 0.25 ml, stored on Eppendorf tubes that were directly introduced on the MR system. Sample preparation For the characterization of the MR properties of magnetoliposomes, we first prepared stock solutions for each of the 4 synthesized liposomal compositions. Since (gadolinium-phospholipids/total amount of phospholipids) ratios were different for the 4 liposomal compositions (x = 0.24, x = 0.12, x = 0.06 and
x = 0.03, for gadolinium-doped phospholipids, giving relative gadolinium amounts of 8x, 4x, 2x and 1x), stock solutions were prepared with respective 1x, 2x, 4x and 8x concentrations of liposomes (i.e. 5, 10, 20 and 40 μmol lipids/ml), in order to achieve the same absolute amount of gadolinium in solution ([Gd]total = 90 μM), as schematically represented in Figure 1. Then, each stock solution was used to prepare 12 seriated dilutions, by using the dilution-extraction method. 24 This method consists on the dilution of the stock solution with a small volume of solvent, retiring later exactly the same volume from the diluted solution, keeping the total volume of the stock solution constant after each dilution-extraction step, and using the retired volume of solution as sample for measurement. Thus, after a series of successive dilution-extraction steps, a precise series of sample solutions of decreasing concentrations is obtained, using low volumes of stock solution (which is quite important when using expensive products). In our case we used a volume of V = 1 ml of stock solution for each liposomal composition to prepare the complete series of solutions (12 solutions for each liposomal composition) to be measured. The retired volume on each dilution step was v = 0.25 ml (Vf factor of 0.8), 24 stored on Eppendorf tubes for MRI studies. A final in vivo study was conducted to assess the sensitivity of detection of the prepared liposomal compositions in the brain. For this study 4 rats were injected with v = 10 μl of liposome solutions with liposomal concentrations of 6.25, 12.5, 25 and
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Figure 3. Color-coded T1 and T2 maps of 12 solutions with variable concentration of liposomes (gadolinium content on solution ranging 10–120 μM) obtained at 9.4 Tesla for the composition of liposomes with the highest payload of gadolinium per liposome unit (white numbers in the background represent calculated concentration of gadolinium in each tube in mM). Similar maps where obtained for the other 3 compositions of liposomes.
50 μM, respectively, for liposomal compositions containing 8x, 4X, 2x and 1x relative amounts of gadolinium lipids. Thus, the total amount of gadolinium injected on each animal was the same, Gdtotal = 15 nanomol (v = 10 μl, [Gd]total = 1.5 mM) . An additional rat was injected with v = 10 μl of a solution 50 mM of Gd-DTPA (Dotarem, Geubert, France).
Results The structure of synthesized liposomes was inspected by acquiring cryo-TEM images of frozen solutions of liposomes (Figure 2, A). These pictures show the formation of (mainly) small unilamellar vesicles (SUVs) with a narrow distribution of sizes in solution of 130 nm, as measured by dynamic lightscattering. A typical histogram of hydrodynamic diameters obtained for liposomal formulations (centered at 129.8 nm) is presented in Figure 2, B (similar results were obtained for other compositions tested). Distributions of sizes present a half-width at half-height of circa 25 nm (24.2 nm in Figure 2, B). Liposomes remain stable in solution for a week, when stored at 4 °C (less than 7% of larger populations detected by DLS at day 7 of storage, data not shown), which is a relevant logistic issue when seriated in vivo studies are planned. For the obtaining of T1 relaxation times, samples were prepared by the dilution-extraction method, as described above, and placed on the MR system on groups of n = 6 samples. An example of T1 and T2 parametric maps obtained for a complete series of solutions of one of the liposomal compositions (x = 0.24, relative 8x load of gadolinium) is presented in Figure 3. On this color-coded image, each circle represents a transversal cut through the sample tube, along with the calculated total concentration of gadolinium. Similar results were obtained for the other three liposomal compositions. Finally, relaxation rates of each solution (R1 = 1/T1) were calculated, and plotted against the total concentration of gadolinium for each solutions, obtaining the linear plots presented in Figure 4. The slope of
each of these linear plots represents the longitudinal relaxivity (r1) of the corresponding liposomal composition. The longitudinal relaxivities obtained were r1 = 2.3, 3.4, 3.5 and 4.9 s - 1 mM - 1 for liposomes containing relative 8x, 4x, 2x and 1x molar fractions of Gadolinium-lipids, which are in the order of most T1 commercial contrast agents for clinical use. 18 We have found a value of r1 = 3.60 ± 0.02 mM - 1 s - 1 at 9.4 T (Figure 4) for commercial Gd-DTPA (DOTAREM), which is in agreement with previously published values for this agent at this magnetic field. 25 Interestingly, the liposomal composition with the lowest molar fraction of Gadolinium-lipids on our study (x = 0.03, relative 1X) presented the highest value for the longitudinal relaxivity. For the in vivo study, Proton density images (with certain T2 weight, at the injection site, introduced by the contrast agent) and R1 maps where obtained for each animal (Figure 5) showing that while the injection of the liposomal composition with the highest payload of gadolinium (molar fraction of x = 0.24) is barely distinguishable form the background signal of brain tissue, the composition including the lowest molar fraction of the tested ones (x = 0.03) is clearly visible, and the contrast induced is comparable to that obtained by a 30 times higher dose of commercial Gd-DTPA. It is noteworthy that solutions of liposomes remain closely contained around the injection site (the mean measured volume of brain parenchyma with reduced T1 time for all animals was 12.4 ± 3.9 μl, after injection of 10 μl of contrast agent), conversely to what happens with the injection of free Gd-DTPA, that diffuses through the brain parenchyma (Vreduced T1 = 71.6 μl, Figure 5). This fact may indicate that liposomes do not behave as a simple solutions (as free Gd-DTPA does) and diffusion along the tissue is restrained, to some extent.
Discussion As we have already mentioned, stealth cholesterol/DSPC liposomes were used as structural basis of our theranostic agents.
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Figure 4. Plot of measured longitudinal relaxation rates (R1 = 1/T1) obtained from T1 maps vs. the total concentration of gadolinium on those solutions ([Gd]Total), for Gd-DTPA and for all studied liposomal compositions. The slope of each of these plots corresponds to the relaxivities (r1) of Gd-DTPA and of each liposomes composition set.
We included a fraction of PEG-conjugated lipids to avoid in vivo aggregation, opsonization and rapid clearance from blood by the reticuloendothelial system. 6,26–29 The development of Stealth technology, by conjugation of polyethylene glycol chains to the surface of liposomes, has been proven to extend the circulatory half-life of the molecules for up to 18 h after systemic injection, increasing their biodisponibility in vivo. 19 Indeed, the integrity and biodisponibility of liposomes, once injected in a living bean, is an important issue to consider for potential in vivo applications of liposomes. Liposomes have been described as one of the most stable systems for drug delivery, 30 and it is known that their integrity in vivo depends on their composition and size. Cholesterol-stabilized liposomes with sizes on the range from 70 to 200 nm (like those prepared here) have been shown to remain in the body for periods of time up to 24–48 h after injection. 30 Some authors have even reported residence times as prolonged as 72 h, assuming that liposomes stored in the spleen of the animals, and released into the blood stream afterwards. 25 Thus the preparation of liposomes like those described in this study, with a mean size of 100 nm, PEG grafts on their surface, and stabilized by cholesterol (our DSPC-Cholesterol liposomes preset a critical temperature over 65 °C) should ensure a sufficient residence of these systems in vivo, to achieve a theranostic activity, as foreseen for these molecules. In addition, the limitation of liposomes’ size to circa 100 nm is expected to facilitate their crossing through the blood brain-barrier, 15,31 which becomes relevant when the molecular event to be followed takes place inside the brain parenchyma, such is the case for theranostic strategies on neurological disorders. Following a multimodal approach, 16,17 phospholipids containing rhodamine or gadolinium were both included into liposomes, to make them traceable by fluorescence and MRI techniques, respectively, thus enabling their use as diagnostic tools. The total load of gadolinium per liposome unit is a critical parameter that has to be optimized for theranostic (molecular
imaging) purposes, as it is crucial for the induction of T1 contrast on MR images. Here, we show that the intensity of the T1 effect induced by a particular load of gadolinium in liposomes is not linearly dependent on gadolinium content. Thus, a series of compositions need to be prepared and tested in order to find the most appropriate payload of the superparamagnetic agent, for optimized sensitivity on in vivo applications. In this work we describe a quick and precise experimental procedure to achieve this goal. Our results confirm the fact, already pointed by other authors, 19 that gadolinium paramagnetic activity is not an additive property when multiple units of this lanthanide are located in the structure of liposomes, in respect to a situation on which gadolinium atoms are free in solution. The relaxivity of magneto-liposomes does not continuously rise by a limitless increase of gadolinium units in their structure. 17,18,32 In fact, tight packing of high amounts of gadolinium units in a reduced space may partially quench the paramagnetic effect of gadolinium, yielding lower global relaxivities. 33 It is important to remember that T1 contrast is generated when water molecules have access to the coordination spheres of gadolinium atoms in solutions. Tight packaging of Gd chelates in liposomes hampers the access of water molecules to those coordination spheres. On the other hand, the use of a large fraction of gadoliniumcontaining lipids in liposomes may also affect the rigidity and permeability of the phospholipid bilayer of liposomes to water molecules. 18,34 Permeability of the membrane of liposomes plays an important role in the final relaxivity of macromolecular molecules, as it has been described by Ghaghada et al. 19 On the other hand, relaxivity of macromolecules containing multiple units of gadolinium chelates shows a great dependency on molecular mobility, 18 but in our study liposomes are always constructed in the same way, with same sizes and total amount of phospholipids. Therefore it may be expected that molecular motions would be similar for all of them, and the rigidity and permeability of liposomes to water molecules may be the predominant effect here. Changes on the water residence time inside the inner cavity of liposomes may be also playing a key role in the contras generated by the designed molecules, which causes a more important effect on global contrast than the electronic relaxivity of gadolinium atoms per se. 18,19,33 Last but not least, the number of gadolinium ions per liposome does not just condition the longitudinal relaxivity of the contrast agent (r1, in mM - 1 s - 1) but does also affect to the transversal relaxation rates (R2 = 1/T2) of water molecules, introducing a negative contrast on T2 weighted MR images. Thus, the contrast achieved on T1 weighted images (which do intrinsically have some T2 weight too) 23 is somehow governed by both relaxivities (r1 and r2), rather than by r1 alone. In fact, an excessive payload of gadolinium may be contraindicated, due to a disproportionate weight of T2 effects in the total contrast. 32 The weight of this effect will be highly dependent on the composition of the system. Thus, Courant et al. 35 have described gadolinium-containing hydrogels with r2 values 2.5 time higher than r1 values (at 1.5 T), claiming a powerful dual T1-T2 effect for those systems. Our liposomes presented a mean r2 relaxivity of 8.38 mM - 1 s - 1, at 9.4 T, which represents 2–3 times higher than observed r1 values, therefore T2 weight is expected to play a
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Figure 5. Grey-scaled Proton density MR images (with certain T2 weight, at the injection site, introduced by the contrast agent) and color-coded R1 parametric maps (with magnification of the injection site) of rat brains obtained in vivo, 1 h after intraparenchymal injections of 500 nanomol of Gd-DTPA (one animal) in 10 μl, of four different formulations of magnetoliposomes, containing the same total amount of gadolinium in solution (15 nanomol), but different lipid concentrations (6.25, 12.5, 25 and 50 μM for 8x, 4X, 2x and 1x liposomes composition sets). The lowest payload of gadolinium units per liposomes (liposomes 1x) induced the highest T1 contrast, in vivo, which resulted comparable to Gd-DTPA injections, even using more than 30 times lower amounts of gadolinium in the liposomes respect to the free agent.
similar role on total contrast as that described by Courant et al. 35 In summary, many are the factors that influence the relaxivity of complex macromolecules, and relaxivity must be optimized case by case. 18 A quick method to perform such optimization, like the one presented here, is highly advantageous. The reduction of the total number of gadolinium units by such a high factor, in order to produce a similar contrast in vivo, is very important to minimize toxic effects from gadolinium nuclei. Gadolinium is a very toxic substance in its free state. For this reason, chelation of Gd units by a coligand (such as DTPA) is required for in vivo applications. 18 This process, on the other hand, reduces the number of coordination sites for water (hydration number, q), and therefore the relaxivity of these molecules. Attempts to increase sensitivity of detection by increasing the number of coordination sites will induce weaker links between gadolinium atoms and the chelating agents, increasing the risk of release of free gadolinium into tissues. Therefore, approaches like packing Gd-chelates in liposomes to achieve contrast are preferable, since Gadolinium lipids keep the highly stable chelation of Gd nuclei by diethylenetriamine pentaacetic acid (DTPA) molecules. Additionally, the low molecular weighted contrast agent (GdDTPA) diffuses rapidly through the brain tissue, and may be easily dumped into one of the lateral ventricles of the brain (Figure 5), while liposomes stay concentrated at the injection site. Free diffusion of the contrast agent through the brain tissue is not a desirable effect for a theranostic agent, which is supposed
to remain at the targeted site, in molecular recognition processes. In this sense liposomal based contrast agents offer superior specificity for in vivo applications in the brain. In summary, in this work we have demonstrated that the intensity of the T1 effect induced by a particular load of gadolinium in theranostic (molecular imaging) molecules may be difficult to predict (complex models may be required for that purpose), 25 and just assuming that liposome’s relaxivities would continuously rise by increasing gadolinium units in their structure may have negative consequences, due to the particularities of the physical mechanisms behind T1 contrast induced by lanthanides. The access of water molecules to the coordination spheres of these molecules is crucial, and highly influenced by the number and disposition of gadolinium units on the structure of macromolecules. We have presented a simple, quick and accurate experimental procedure to estimate the relaxivities of MR theranostic nano-platforms based on gadolinium, using minimal amounts of these expensive molecules. Results were confirmed by in vivo studies in the rat brain.
Acknowledgments Authors acknowledge Prof. Dr. Klaas Nicolay and Dr. Gustav Strijkers, From Eindhoven University of Technology (The Netherlands), for sharing their knowledge on the synthesis and characterization of liposomes.
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