Multifunctional Magnetoliposomes for Sequential ... - ACS Publications

Multifunctional Magnetoliposomes for Sequential Controlled Release Annalisa Salvatore, Costanza Montis, Debora Berti,* and Piero Baglioni Department of Chemistry and CSGI, University of Florence, Via della Lastruccia 3, 50019-Sesto Fiorentino, Florence, Italy S Supporting Information *

ABSTRACT: The simultaneous or sequential delivery of multiple therapeutic active principles to a specific target is one of the main challenges of nanomedicine. This goal requires the construction of complex devices often extremely time and cost consuming. Supramolecular selfassemblies, with building blocks of different nature, each providing a specific function to the final construct, can combine a facile synthetic route with a high tunability and structural control. In this study we provide the proof-ofprinciple of a drug delivery system, DDS, constituted of (i) liposomes, providing a fully biocompatible lipid scaffold suitable to host both hydrophobic and hydrophilic drugs; (ii) a double-stranded DNA conjugated with a cholesteryl unit that spontaneously inserts into the lipid membrane; and (iii) hydrophobic and hydrophilic superparamagnetic iron oxide nanoparticles (SPIONs) embedded inside the lipid membrane of liposomes or connected to the DNA, respectively. Upon application of an alternating magnetic field, the SPIONs can trigger, through thermal activation, the release of a DNA strand or of the liposomal payload, depending on the frequency and the application time of the field, as proved by both steady-state and time-resolved fluorescence studies. This feature is due to the different localization of the two kinds of SPIONS within the construct and demonstrates the feasibility of a multifunctional DDS, built up from self-assembly of biocompatible building blocks. KEYWORDS: magnetoliposomes, SPIONs, DNA delivery, controlled release, antisense oligonucleotides, core−shell NPs, fluorescence correlation spectroscopy

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molecular recognition abilities toward biological targets, and thermal responsiveness. The combination of these features with the optical, thermal, and electric properties of inorganic materials can provide fine control over material characteristics.6 Both lipid assemblies and DNA are thermoresponsive. In the first case, below the melting temperature of lipid bilayers (Tm), lipids are in the rigid, viscous, and highly organized gel phase, while above Tm the lipids chains are in a liquid crystalline fluid phase. Their permeability is enhanced at Tm. By choosing the lipid composition, it is possible to tune the transition temperature and bilayer permeability and, thus, control the release of drugs encapsulated in the aqueous pool. The possibility to remotely induce this phase transition and control drug release through external magnetic fields has sparked the interest in magnetoliposomes,7 where superparamagnetic nanoparticles (SPIONs)8 located in the bilayer,9−12 in the internal pool of liposomes,13,14 or bound to the surface15 generate local temperature gradients due to their magnetic responsiveness.16 Over the past years, numerous examples of

anomedicine holds the promise to be the main enabling technology for earlier and more precise individual diagnosis and ground-breaking targeted therapies.1 The possibility to selectively reach a biological target with limited side effects, high efficiency, biocompatibility, tailored pharmacokinetic profile, and space, time and concentration control of the therapeutic/diagnostic principles inside the body is one of the most important challenges in the field.2 Self-assembly is an ingenious, cost-effective, and convenient strategy to build up complex architectures, where the synergy of weak forces, as van der Waals, π-stacking, and H-bond interactions, leads to fine control of size, shape, and function. Self-assemblies are responsive to slight variations of control parameters, such as pressure, temperature, and ionic strength, and seemingly subtle changes can produce cascade effects from the molecular scale to the mesoscale. Although chemically very different, lipids and oligonucleotides (ONs) are among the most attractive building blocks for nanotechnology. While lipid assemblies, safe and FDA approved,3 can be employed as scaffolds for multifunctional devices, the versatility and specificity of DNA as a modular building block for nanotechnology4,5 lie in its selectivity, © 2016 American Chemical Society

Received: May 13, 2016 Accepted: August 9, 2016 Published: August 9, 2016 7749

DOI: 10.1021/acsnano.6b03194 ACS Nano 2016, 10, 7749−7760

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ACS Nano magnetoliposomal devices have been reported, principally for imaging,17 biosensing,18 or heating by using lasers,19 alternating current electromagnetic fields10 (AC EMFs), or alternating magnetic fields11,20−22 (AMFs). Our group has studied hybrid lipid/SPION architectures for controlled drug release. In particular, we introduced the concept that a low-frequency alternating magnetic field (LF-AMF), below the hyperthermic threshold, produces a release of the payload from giant unilamellar vesicles,23 liposomes,14,24,25 or cubosomes26 loaded with SPIONs, due to a local temperature increase that enhances the permeability of the lipid membrane. Concerning DNA, with a similar principle, the conjugation of SPIONs to ONs can be applied to control strand dissociation of DNA through application of a magnetic field. We have recently shown the formation of thermoresponsive SPION clusters connected through DNA double strands, where the application of LF-AMF causes the release of a single-stranded (ss)-ON and the disruption of the cluster.27 Nucleic acids,28 ssDNA and RNA, antisense ONs,29,30 and aptamers are among the most promising systems in the development of specific therapies for a plethora of diseases,31 but their targeted delivery to cells still poses severe challenges, because of the negligible permeability of cell membranes and of their susceptibility to degradation by nucleases. The combined delivery of nucleic-acid-based therapeutics with additional drugs can be particularly attractive to achieve a synergistic effect of the active principles. For instance, plasmid DNA for gene therapy has been co-delivered both with an antiinflammatory for cardiac32 and bowel33 diseases and with chemotherapeutics for cancer treatments.34 Here we present a lipid/DNA/SPION nanostructured architecture, designed as a multifunctional drug delivery system able to transport a therapeutic oligonucleotide, which is in particular an antisense sequence with down-regulation properties toward the receptor ERB B2, involved in diabetic peripheral neuropathy (DPN),35 which can be co-delivered with an additional anti-inflammatory principle, to address both the causal and symptomatic features of the disease.36 Scheme 1 summarizes the construction strategy. 1,2Dipalmitoyl-sn-glycero-3-phosphocholine (DPPC) liposomes (a) are decorated with two different kinds of magnetic nanoparticles: hydrophobic Fe3O4 NPs, embedded in the lipid bilayer (b), and hydrophilic Fe3O4 NP cores coated with a gold shell (Au@Fe3O4 NPs). The Au shell is functionalized with a ss-ON hybridized with a DNA therapeutic zipper and with a cholesteryl-ON, which inserts in the lipid membrane (c). The final device (d) is easily built up through spontaneous assembly of the different building blocks, programmed to interact through a predetermined pathway. The construct defines two separated hydrophilic compartments. The first one is the liposomal aqueous corona, where we locate the DNA zipper, i.e., ERB B2 antisense oligonucleotide, as a model ONbased drug, and the second one is the liposomes’ lumen, where we confine a fluorescent molecule as a model hydrophilic drug to be co-delivered. In order to control the release of the model therapeutics, we inserted two different magnetic triggers, a hydrophobic and hydrophilic SPION, located in the liposomal aqueous corona and lipid membrane, respectively. The local temperature gradients induced by AMFs affect the permeability of the membrane and the stability of the DNA zipper and can be tuned to control simultaneous or sequential release of the two principles.

Scheme 1. Sketch of the DDS design principle: DPPC liposomes (L, a) are loaded with hydrophobic magnetite nanoparticles (black spheres) coated with oleic acid and oleylamine, to obtain magnetoliposomes (ML, b); both liposomes (c) and magnetoliposomes (d) are then decorated with a single-stranded ON (pink) tagged with a cholesteryl unit that inserts into the lipid membrane, connected to another single-stranded ON (light blue) conjugated to hydrophilic magnetite nanoparticles coated with a gold shell, through a complementary single-stranded therapeutically active DNA zipper (purple)

In the following sections, we will report and discuss the stepby-step construction of the device, its structural characterization, and its responsiveness to different low-frequency alternating magnetic fields.

RESULTS AND DISCUSSION Preparation and Characterization of the Nanocarrier. The controlled release is exerted through two thermally activated processes occurring at different temperatures, i.e., the permeability increase of the lipid membrane at the Lβ−Lα transition (41 °C) and the double-stranded (ds)-DNA melting with release of a ss-ON (53 °C). This feature opens up the possibility of using the same magnetic nanoparticles, whose different surface coatings (hydrophobic and hydrophilic) promote spontaneous insertion at different sites of the device, to trigger the selective release of the desired active principles. The first building blocks of the nanocarrier (Scheme 1d) are the two different types of magnetic NPs, i.e., hydrophilic SPIONs, conjugated to a DNA zipper, and hydrophobic SPIONs embedded in the lipid membrane. Their response to AMF causes a local temperature increase, due to hysteresis losses or Brownian and Néel relaxations, which depend on several parameters, such as size, surface coating (e.g., with a metal shell), viscosity of the medium, and AMF frequency.37−39 In addition, the presence of a Au shell and the localization in the aqueous corona of the liposomes or within the lipid membrane, respectively, which differ in viscosity by about 2 orders of magnitude,40 can substantially modulate this process. Hydrophobic and hydrophilic SPIONs were prepared adapting the procedure reported by Wang41 for the synthesis of Fe3O4 cores, as described in the Methods section. This synthesis leads to 5 nm NPs (see TEM and SAXS, Figure 1a), 7750


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Figure 1. Representative TEM images and SAXS profiles of (a) SPIONs in hexane and (b) core−shell NPs in water after ligand exchange. The SAXS curves have been fitted according to a spherical form factor with Schulz polydispersity and a polydisperse core−shell model for the core/shell NPs (see the Methods section). (c) UV−vis spectrum of an NP-chol-TEG-DNA before (dotted line) and after (full line) hybridization, and melting curve of the DNA-decorated NPs measured at 260 nm, with heating (1) and cooling (2) ramps (inset). SAXS spectrum of iron oxide NPs in hexane (a) is reprinted from ref 26 with permission.

capped with a hydrophobic coating of oleylamine and oleic acid, which enables affinity to the lipid bilayer, while the small size minimizes disruption of the membrane integrity. Hydrophilic Au@Fe3O4 NPs were synthesized by coating the hydrophobic ones, according to the procedure described in the Methods section. The Au coating enhances biocompatibility, protects the core against oxidation, and provides a wellestablished platform for chemical functionalization. In our case we conjugated a thiolated ON with half of its strand complementary to an antisense ON (zipper DNA). 42 Conjugation of the thiolated ON with Au@Fe3O4 was performed through ligand exchange with a 1:1 mixture of methoxy-PEG thiol and thiolated ON, acting both as phase transfer agents for the nanoparticles (from hexane to water) and as capping agents of the Au@Fe3O4 in water. In addition, the PEG coating with size comparable to the length of the oligonucleotide provides a valuable protection to the genetic material, preventing its degradation.43 Figure 1b reports a representative TEM image and SAXS curve of the core−shell NPs, after the ligand exchange. The average diameter of Au@ Fe3O4 NPs can be estimated as 7 ± 1 nm (from TEM and SAXS), with a gold shell thickness of about 1 nm, as determined from the SAXS analysis (see SI for details). Finally, the ss-ON conjugated to Au@Fe3O4 NPs was annealed with the complementary DNA sequence (a 1:1 mixture of the DNA zipper and of the cholesteryl-ON) to achieve the strand pairing. The UV−vis absorbance spectra of NP conjugated with DNA before (dotted line) and after (solid line) annealing are reported in Figure 1c. The NP scattering sums up to the DNA absorbance (260 nm) and Au plasmon resonance (520 nm), further confirming the successful attachment of the thiolated ON. DNA melting and hybridization can be followed with a thermal scan of the absorbance at 260 nm (inset, Figure 1c). The molar absorptivity increases upon DNA melting and

symmetrically decreases upon cooling the sample to room temperature, highlighting the reversible nature of the process. Liposomes (L) and magnetoliposomes (ML) were prepared through lipid film hydration coupled with sequential extrusion, introducing some modifications to include hydrophobic nanoparticles (see Methods section). The density of NPs decorating the lipid membrane is chosen as a balance between membrane stability when no AMF is applied and efficient responsiveness under AMF. High NP/lipid ratios do not strictly translate into larger release and could compromise membrane integrity and lead to passive leakage.9,10,24 In the present case, the NP/lipid mass ratio, evaluated through ICP-AES, is 0.011, which is an intermediate concentration, as compared to the examples reported in the literature,10,11 and roughly corresponds to 10 NPs/liposome. In order to shed light on the structure and thermotropic properties of L and ML, we performed SAXS-WAXS experiments on DPPC multilamellar vesicles (MLVs), i.e., liposomal dispersions before extrusion, and unilamellar vesicles (ULVs), in the absence and in the presence of SPIONs. Figure 2 displays the SAXS and WAXS curves obtained for DPPC MLVs. Wide-angle X-ray diffraction (WAXD) patterns at 25 °C show the presence of a sharp reflection at 1.49 Å−1 with a broad shoulder around 1.51 Å−1 (Figure 2b), typical of lipid bilayers in the lamellar gel phase with tilted hydrocarbon chains and pseudohexagonal chain packing.44 Upon temperature increase, WAXD is consistent with the melting of the phospholipid chains, with a pretransition between 30 and 35 °C and the main transition between 40 and 45 °C. The SAXS profile is characterized by the presence of Bragg reflections due to the periodicity of the lamellar stacks. The smectic period (d) of the lamellar phase, 6.4 nm at 25 °C, increases to 6.9 nm above the Lβ−Lα transition temperature (Tm = 41 °C, see Figure 2c). It is well known that the phase transition causes the loss of the tilt angle (that would increase the d-value by about 0.55 nm) and a simultaneous decrease of d of about 0.8 nm due to 7751


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Figure 2. (a) SAXS and (b) WAXS patterns of DPPC multilamellar vesicles (MLV, 100 mg/mL with (DPPC-ML) and without (DPPC-L) NPs embedded in the lipid bilayer, from 25 to 50 °C; (c) temperature trend of the smectic period of DPPC lamellar phase (d) in the absence (filled circles) and in the presence (empty triangles) of NPs; (d) representative SAXS experimental curves (empty circles) and fitting curves (solid lines) of unilamellar (ULV, 2.5 mg/mL) magnetoliposomes (yellow) and liposomes (gray) acquired at 25 °C; (e) scheme sketching the bilayer parameters as obtained from SAXS data: smectic period of the lamellar phase (d), bilayer thickness (dB), thickness of the water layer separating the lamellar stacks (dW); (f) temperature variations of the bilayer parameters: bilayer (dB) and water thicknesses (dW) for liposomes (L, circles) and magnetoliposomes (ML, triangles).

lamellar phases, while for ML, a contribution of the nanoparticles to the scattering intensity was also included14 (see the Methods section). The temperature trend of the structural parameters is reported in Figure 2f. As expected, the bilayer thickness (dB) of both L and ML decreases, due to chain melting, although this effect is stronger in the absence of NPs. The insertion of the nanoparticles causes a deformation of the lipid membrane, detected as a slight increase in the bilayer thickness at room temperature. As we have seen, for neat MLVs the temperature increase determines an increase of the interlamellar water layer, while this effect is reversed in the presence of nanoparticles, probably due to their hydrophobic nature. Finally, the presence of the NPs clearly decreases the order within the bilayer, determining the loss of the gel−liquid crystalline pretransition at 35 °C. However, NPs do not modify the overall bilayer integrity, proving that, similarly to L, ML are able to host hydrophilic therapeutic species and retain them inside the aqueous pool. Figure 3a compares the dynamic light scattering (DLS) profiles obtained for liposomes (black circles) and magnetoliposomes (red triangles). The hydrodynamic diameter,

conformational variations of the acyl chains upon melting. Therefore, the overall increase of d is due to the larger thickness of the interlamellar water layer.45 In the presence of NPs (Figure 2a, DPPC-ML), the lamellar spacing at 25 °C increases (7.6 nm), confirming the presence of the NPs embedded in the lipid membrane. Moreover, the NPs modify the thermal behavior of the bilayers. While the smectic period of MLVs without NPs shows a sharp increase around Tm (see Figure 2c), in the presence of SPIONs the structural parameters show a smoother and opposite trend, leading to an overall continuous decrease of d. Furthermore, the sharp reflections in WAXS (Figure 2b) observed for pure DPPC MLVs are replaced by a quasi-symmetric scattering around 1.5 Å−1, indicating that the hydrocarbon chains are oriented perpendicularly to the bilayer plane, in agreement with a zero tilt angle of the acyl chains.46 This change in orientation of the phospholipid chains also explains the disappearance of the pretransition between 30 and 35 °C. SAXS was also performed on ULV in the absence (L) and in the presence (ML) of SPIONs (Figure 2d). In the first case we analyzed the spectra according to the Nallet model47 for 7752


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Figure 3. (Upper panel) Representative normalized dynamic light scattering (DLS) curves obtained for (a) liposomes (black circles) and magnetoliposomes (red triangles) and (b) for the same samples before (empty markers) and after (filled markers) insertion of the NP-dsDNA unit: liposomes (black circles, L and NP-dsDNA@L, respectively) and magnetoliposomes (red triangles, ML and NP-dsDNA@ML, respectively). The ACFs were acquired at a 0.03 mg/mL lipid concentration. (Lower panel) Representative fluorescence correlation spectroscopy (FCS) of the emission of a fluorescent probe that selectively binds to ON: (c) Oligreen bound to the zipper-DNA (black filled markers) and to NP-dsDNA (orange empty markers); (d) Oligreen bound to the NP-dsDNA unit inserted inside liposomes (black circles) and magnetoliposomes (blue triangles); the continuous lines in (c) and (d) indicate the best fit according to a 3D one-component diffusion model (see Methods section).

obtained through a second-order cumulant fit, is 90 ± 15 nm for liposomes and 150 ± 30 nm for magnetoliposomes. We can ascribe this size increase to a possible clusterization of the ML promoted by the embedded NPs. Thus, the distribution of DNA inside ML could be slightly different than in L, due to steric effects caused by ML clusterization. Previous studies reported that clustering of embedded NPs is observed when the concentration of the NPs is relatively high,10 whereas NPs with diameters from 2 to 6.5 nm can reside in bilayer compartments within individual or neighboring liposomes, partially promoting the association of liposomes themselves48,49 when the lipid:NP ratio is high. Overall, DLS data revealed the successful spontaneous encapsulation of the hydrophobic nanoparticles in the lipid membranes and the formation of magnetoliposomes. We then added the NP-dsDNA block, which spontaneously insert in L or ML due to its cholesteryl tag (see Scheme 1d). An aliquot of the NP-dsDNA dispersion was added to L or ML and left at room temperature under stirring for 24 h (see Methods section). Figure 3b reports the main DLS results: regarding liposomes, the hydrodynamic diameter significantly increases from 90 ± 8 nm to 120 ± 10 nm with the addition of the NPdsDNA, confirming the successful conjugation of the lipid architecture with the NP-dsDNA unit. On the other hand, the conjugation of the DNA unit on ML does not affect significantly the hydrodynamic diameter, probably due to the polydispersity of the ML themselves that can shadow the additional hydrodynamic thickness caused by insertion of the DNA-NP unit.

In order to confirm the assembly, we monitored the NPdsDNA@L(ML) conjugation through fluorescence correlation spectroscopy (FCS). This technique records the fluctuations of the emission intensity of fluorescent species moving inside and outside the excitation volume of a confocal microscope; such fluctuations are analyzed in terms of the autocorrelation function of the fluorescence intensity (ACF), which contains information on the diffusion properties of the fluorescent species. In particular, the intercept is inversely proportional to the concentration of the probe, the decay time is connected to the diffusion coefficient, thus to the hydrodynamic radius and the viscosity of the medium, and the decay profile is determined by the diffusion mode of the species. Here we investigated the insertion of the NP-dsDNA moiety in L and ML using the intercalating fluorescent tag (Oligreen) that selectively labels ONs. Oligreen is characterized by an increased fluorescence quantum yield when bound to DNA. Figure 3c displays the FCS curves acquired for Oligreen conjugated to the ss-zipper (black filled markers) and for the tag bound to the NP-dsDNA (orange empty markers), respectively. Both curves were analyzed according to 3D Brownian diffusion of a single species (see Methods section for details), leading to diffusion coefficients of D = 112 ± 10 μm2 s−1 for the ON-dye and D = 60 ± 20 μm2 s−1 for the NP-dsDNA, respectively. These diffusion coefficients are consistent with the expected values. In fact, if we consider the NP-dsDNA unit as a rodlike arm, the estimated diffusion coefficient50 for a 19-base length of about 120 μm2 s−1 is decreased by the addition of 6 thymine bases and a 5 nm diameter NP (see Figure 1) to about 60 μm2 s−1, in full agreement with the FCS measurements. After incubation with liposomes (Figure 3d black circles) and magnetoliposomes 7753


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Figure 4. FCS curves from liposomes (a, c) and magnetoliposomes (b, d) at a concentration of 125 μg/mL with chol-DNA-NP inserted, upon application of a low-frequency alternating magnetic field at 3.22 kHz (upper panel) and 6.22 kHz (lower panel) for increasing times (5 min blue curves, 10 min orange curves, 15 min green curves); the reported FCS curves are normalized with respect to that measured at 5 min, for clarity; the FCS were analyzed according to a 3D two-component Brownian diffusion model, as described in the Methods section, leading to the evaluation of the relative fractions of free DNA and DNA bound to the liposomes; (e, f) percentage of ss-DNA released in time as evaluated from FCS upon application of an AMF at (e) 3.22 kHz and (f) 6.22 kHz on magnetoliposomes (ML, green) and liposomes (L, blue).

zipper ON and a model hydrophilic drug (carboxyfluorescein) contained in the aqueous pool. Clearly, the presented multifunctional architecture is complex. The fundamental contributions to local heating effects due to AMF application to SPIONs are the Brownian and Nèel relaxations. While the efficacy of the Nèel contribution depends on the magnetic properties of the NPs (which in turn depend on the core composition and on the size)51 and on the characteristics of the applied magnetic field,37 the Brownian contribution depends on the diffusion properties of the SPIONs (i.e., viscosity of the medium, size, and diffusion constraints). Furthermore, direct interaction of the SPIONs, resulting in dipole−dipole contributions, and contributions due to surface functionalization of the SPIONs (in particular with metal shells) are also effective in the temperature rise. In our system, the simultaneous presence of two types of SPIONs with two different localizations and with both dipole−dipole and gold shell contributions, makes the interpretation of the relaxation mechanisms not straightforward. It is however well established that the effect of the AMF application on SPIONs is a local temperature increase rather than a bulk effect.52−54 To confirm this point, we monitored

(Figure 3d blue triangles) the decay times increase dramatically, highlighting the successful and quantitative insertion of the NPdsDNA in the lipid membranes. The curves were analyzed according to a one-component 3D Brownian diffusion, leading to diffusion coefficient D = 4.3 ± 1.2 μm2 s−1 for NP-dsDNA@ L and D = 1.98 ± 0.55 μm2 s−1 for NP-dsDNA@ML. The diffusion coefficients obtained from the assembled lipid−DNA complexes well agree with those obtained from DLS measurements, demonstrating that the same diffusing species are monitored with the two techniques and that selfassembly indeed occurs, as programmed, through design of the building blocks. Moreover, the fact that FCS data can be analyzed with a single-component diffusion proves that the conjugation of the NP-dsDNA moiety to liposomes is quantitative, since no residual decay from the ds-DNA-NP species is detected. These data unambiguously show that the nanoarchitecture assembles as predicted by design of the building blocks, with fine structural control and quantitative yield. Controlled Release. The responsiveness of the nanocarrier to LF-AMF of different frequencies and application times was tested in terms of controlled release of two model drugs, the 7754


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Figure 5. Percentage of CF leakage during AMF application reported as fluorescence intensity recovery for different application times, (a) for L and ML, (b) in the presence of a DNA unit.

interparticle distance,59 as well as lower the magnetization of the core,41 leading to a higher energy of magnetic anisotropy.56 However, it has been shown that the overall effect of a gold shell is to increase the local temperature rise, probably due to the higher conductivity of Au that enhances the formation of eddy currents on the surface60,61 and, thus, the losses due to this phenomenon. Figure 4e and f report the relative abundance of the released zipper ON, upon application of 3.22 kHz (e) and 6.22 kHz (f) AMF on liposomes (L, blue line) for 5, 10, and 15 min. Even after 15 min, no major differences can be observed for L, with a zipper release of 17% and 15% achieved for 3.22 and 6.22 kHz, respectively. Overall, the efficiency of this device can be deemed as low. Conversely, in magnetoliposomes the zipper release is quantitative for 6.22 kHz after 15 min. This indicates a substantial contribution of the hydrophobic SPIONs to the local achievement of dsDNA melting temperature. The temperature gradient around a hot NP decays with the reciprocal distance.62 By taking into account that one ML is decorated by about 30 NP-dsDNA and 10 hydrophobic Fe3O4 NPs, the average linear distance of a dsDNA unit from a hydrophobic SPION on the ML is about 20−30 nm. Across this distance, the temperature gradient should be considerably damped. However, the distance is reasonably overestimated, since the relative fluidity of the membrane and the mobility of the dsDNA arm presumably allow the hydrophobic NPs and the dsDNA to come into closer proximity, and the synergistic action of the two types of NPs promotes the full release of the DNA zipper. This quantitative on-demand release of a DNA strand is the result of this architecture. In previous studies on similar magnetic NPs in solution, but without liposomal anchoring, the release after 15 min of a 6 kHz AMF was about 30%,27 which poses limits to the applicability in therapy or for nanotechnology purposes. In the present case, the application of a magnetic field controls a quantitative DNA release. The final aim of this work was to achieve sequential or simultaneous release of the DNA therapeutic sequence and of

the bulk temperature increase induced by coil heating effects, in the absence and in the presence of NPs (see SI for details). Figure 4 shows the FCS curves from the nanocarrier containing only hydrophilic SPIONs (a, c) and both hydrophilic and hydrophobic SPIONs (b, d) upon application of LFAMF of two different frequencies (3.22 and 6.22 kHz) for different times (5, 10, or 15 min). The AMF causes a local temperature increase due to the relaxation of the SPIONs, eventually reaching the dsDNA melting temperature, with subsequent release of the therapeutic zipper ON. This is proved by the variation of the FCS profiles, with gradual appearance of a contribution from the released zipper DNA and a concomitant decrease of the intercepts. This latter effect, ascribable to an increase of the apparent concentration of the probe (see SI, Table S8), is actually connected to the dequenching effect that the fluorescent tag experiences once the DNA zipper is released in solution away from the gold shell of Au@Fe3O4 NPs. The FCS profiles after application of the LF-AMF are fully accounted for by a two-component 3D normal diffusion model, which combines the contribution of the tag bound to the released zipper (D2 = 112 μm2 s−1, as determined from the previous FCS experiments) and that of the dye still associated with the cholesteryl-ON inserted into ML or L (with D1 = 1.98 μm2 s−1 for ML and 4.3 μm2 s−1 for L, as determined from the previous FCS experiments). By keeping constant the predetermined diffusion coefficients, the relative contributions of the two populations (the dsDNA decorating liposomes and magnetoliposomes and the free ON) can be determined as the sole fitting parameters of the autocorrelation functions (see the Methods section for details). Concerning liposomes (L, Figure 4a,c), the application of an LF-AMF causes a local temperature increase,24,55,56 due to several loss processes of Au@Fe3O4 NPs, depending upon specific experimental conditions.57,58 The presence of a gold shell is known to modify the responsiveness of the magnetic core to external magnetic fields. It can reduce the dipole coupling between neighboring NPs, due to increased 7755


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ACS Nano another hydrophilic drug retained within the same nanostructure and accomplished with the same kind of external stimulus. As a final step, we turned therefore our attention to the actuated release of a model hydrophilic drug, carboxyfluorescein (CF), encapsulated in the aqueous pool of L or ML. CF was added during hydration of the lipid film at a concentration of 10 mM and removed from the dispersing medium with extensive dialysis after liposome formation. As reported in the literature,63 concentrated CF solutions exhibit a self-quenching effect. The recovery of emission intensity indicates leakage from the liposomes and therefore an increase of membrane permeability upon application of stimuli.10,55 We monitored CF emission intensity as a function of the AMF frequency and application times. At the end of each series of measurements, the liposomes were disrupted by addition of a detergent (Triton X-100, 1% v/v) to provide an absolute scale (100%) for the experiments. Figure 5a displays the results for L and ML at 3.22 and 6.22 kHz, reported as fluorescence intensity percentage for different application times. As expected, the application of a magnetic field does not induce significant leakage if the NPs are not present, and no dependence on field frequency and application time can be appreciated (Figure 5a, right). Conversely, for magnetoliposomes we have, as expected, a clear effect of the magnetic field on the membrane permeability, indicated by the CF leakage, which reaches 100% after 15 min, 3.22 kHz. Nevertheless, it is important to mention that the application of LF-AMF does not lead to the liposomes’ rupture (see SI for details). In the presence of the DNA carrier (Figure 5b) the emission of CF released in solution does not reach the 100% value for both L and ML, even upon longer (15 min) applications of AMF. This seems to contradict the previously discussed synergistic effects of the SPIONs embedded in the lipid bilayer and the Au@Fe3O4 concerning the release of the DNA zipper. However, as previously discussed, the quenching effect of the gold shell of the NPs, which is a well-established and general phenomenon, possibly causes a decrease in the observed fluorescence intensity, with consequent underestimation of the released CF. Interestingly, a significant CF release in solution is observed for NP-dsDNA@L samples (Figure 5b): this can be attributed to proximity of the Au@Fe3O4 NPs to the lipid membrane. The presence of the triethylene glycol (TEG) and oligoT moieties on the NP-dsDNA units imparts a relative mobility to the arm that is able to rotate and interact with the lipid membrane of liposomes and magnetoliposomes. These results prove that both types of SPIONs synergistically contribute to the release of molecules entrapped in the aqueous pool of magnetoliposomes and from the DNA zipper. Overall, the FCS and steady-state fluorescence data prove that we are able to exert a sequential and controlled release of two different drugs from the drug delivery system (DDS), by simply exploiting the different magnetic properties and localization of the two different SPIONs, embedded in the lipid bilayer and connected to the DNA zipper, respectively. This possibility is depicted in Scheme 2. Briefly, upon application of a 3.22 kHz AMF for short times, the release of a drug contained in the aqueous pool occurs: in fact, the efficiency of the actuated release is due to the temperature gradient caused by the relaxation of the hydrophobic SPIONs inside the lipid bilayer, which triggers the gel to liquid

Scheme 2. Representative scheme of the DDS functionality for sequential controlled release: the application of a 3.22 kHz AMF for short times is sufficient to provoke the release of the hydrophilic drug contained in the aqueous pool of magnetoliposomes. Subsequently, the application of a 6.22 kHz AMF for longer times allows reaching the DNA melting temperature and the release of the zipper therapeutic ON.

crystalline transition (41 °C) and promotes the permeability increase of the membrane. Subsequently, the application of a slightly higher AMF (6.22 kHz) for longer times is required to achieve the DNA melting temperature (53 °C), due to the synergistic effect of the hydrophobic and hydrophilic SPIONs, both contributing to the melting of the DNA zipper. Overall, this architecture allows for a spatial−temporal-controlled release of therapeutics: by varying the nature of the SPIONs and their concentration, it is possible to build up smart nanovectors with a broad applicative field.

CONCLUSIONS We have successfully prepared a complex architecture, through simple self-assembly steps, containing two different carriers, a liposomal scaffold and a DNA zipper, and two types of magnetic nanoparticles as triggers for the magnetically actuated release. The major advantage of this architecture is in the possibility to sequentially control the release of different cargoes by varying the frequency of the field applied. We have shown that 5 nm SPION NPs can be embedded into DPPC bilayers to control the release of a model drug from the liposomal interior with an LF-AMF. We have reached a high percentage of drug release even at low frequencies (Figure 5a), and the control experiments have proved that the spontaneous leakage, influenced by the warming of the dispersion generated by the alternating electric current, is negligible. Moreover, through a DNA-intercalating fluorescent probe we have proved the quantitative conjugation of the NP-DNA unit to the magnetoliposomes, which occurs through spontaneous insertion of the cholesteryl tag in the lipid bilayer. By monitoring the diffusion coefficient of the probe we have evaluated that both hydrophobic SPIONs embedded in the bilayer and the hydrophilic NPs attached on the surface are involved in the release of the payload encapsulated within the liposomal aqueous pool at LF-AMF. In turn, the increase in the apparent concentration of the fluorescent probe in solution, as a consequence of the decreased quenching effect of gold, is an additional evidence that the DNA is released by LF-AMF, leading to a very high efficiency with respect to previous studies. Furthermore, the presence of SPIONs in the liposome bilayers should allow steering the vesicles to the desired site of action using static magnetic fields, providing therefore means for remote targeting. Additional experiments are needed to determine the stability and leakage characteristics in serum and in vivo. 7756


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etry (ICP-AES), and the concentration of the Fe3O4 NPs obtained was 0.011 mg/mL. UV−Vis Spectroscopy. UV−vis absorption was measured on a Cary UV 100 spectrophotometer (Agilent Technologies, Italy) equipped with a 6 × 6 temperature controller for melting analyses. Annealing of the complementary sequences was registered at 260 nm in the range between 20 and 80 °C with a temperature ramping of 0.5° min−1. The acquired spectra were measured for a 3 μg/mL ON solution and for the same sample with an equimolar concentration of complementary ss-DNA zipper and ssDNA conjugated to the core− shell NPs. Small-Angle and Wide-Angle X-ray Scattering. SAXS measurements were carried out on an S3-MICRO/instrument (HECUS, GmbH, Graz, Austria), which consists of a GeniX microfocus X-ray sealed Cu Kα source (power 50 W), which provides a detector-focused X-ray beam with a 0.1542 nm Cu Kα line. The instrument is equipped with a beam with two one-dimensional (1D) position-sensitive detectors (HECUS 1D-PSD-50 M system). Each detector is 50 nm long (spatial resolution 54 um/channel, 1024 channels) and covers the SAXS Q-range (0.003 < Q < 0.6 Å−1) and WAXS Q-range (1.2 < Q < 1.9 Å−1). The temperature was controlled by means of a Peltier TCCS-3 Hecus. SAXS curves were recorded from 25 to 50 °C in a glass capillary for samples in liposomal phase and in a solid sample holder for the multilamellar phase. A lipid concentration of 100 mg/mL was employed for multilamellar vesicles and of 2.5 mg/mL for unilamellar vesicles with an amount of magnetite nanoparticles equal to 1.5 mg/mL and 40 μg/mL for multilamellar and unilamellar vesicles, respectively. Transmission Electron Microscope. Trasmission electron microscopy (TEM) images were acquired with a STEM CM12 Philips electron microscope. The nanoparticle samples dispersed in hexane and in water were cast onto a carbon-coated copper grid sample holder, followed by evaporation of the solvent at room temperature. Dynamic Light Scattering. DLS measurements were carried out on a Brookhaven Instruments apparatus (BI 9000AT correlator and BI 200 SM goniometer). The signal was detected by an EMI 9863B/350 photomultiplier. The light source was the second harmonic of a diode Nd:YAG laser, λ = 532 nm, Coherent DPY315M-100, linearly polarized in the vertical direction. The normalized electrical field time autocorrelation of the scattered light was measured at 90 and analyzed according to the Siegert relationship (eq 1), which connects the firstorder or field-normalized autocorrelation function g1(q, τ) to the measured normalized autocorrelation function g2(q, τ):

METHODS Materials. Fe(III)acetylacetonate, 1,2-hexadecanediol, oleylamine, oleic acid, diphenyl ether, ethanol, and hexane employed for the synthesis of the NPs, and the carboxyfluorescein used as a model drug, were purchased from Sigma-Aldrich (St. Louis, MO, USA). Au(acetylacetonate) was purchased from Alfa Aesar. All the ON sequences were purchased from ATDbio Ltd. (Southampton, UK). The phospholipid was purchased from Avanti Polar Lipids, and the Oligreen fluorescent dye from Invitrogen (Heidelberg, Germany). Synthesis of Magnetic Nanoparticles. Iron oxide nanoparticles were synthesized according to the procedure reported by Wang et al.41 Briefly, 0.71 g of Fe(acac3)3 (2 mmol) was dissolved in 20 mL of phenyl ether with 2 mL of oleic acid (6 mmol) and oleylamine (4 mmol) under a nitrogen atmosphere and vigorous stirring. 1,2Hexadecandiol (2.58 g, 10 mmol) was added into the solution. The solution was heated to 210 °C, refluxed for 2 h, and then cooled to room temperature. Ethanol was added to the solution, and the precipitate collected, washed with ethanol, and dispersed in 20 mL of hexane. A stable dispersion of the magnetic NPs with a hydrophobic coating of oleic acid and oleylamine in hexane was thus obtained. Synthesis of Core−Shell NPs with ds-DNA. For the preparation of the nanoplatform with the delivered DNA, we used the iron oxide NPs previously reported before the purification with ethanol: after refluxing for 2 h and cooling to room temperature, the reaction solution was directly used without separation. In a typical synthesis, 10 mL of the phenyl ether reaction solution of Fe3O4 nanoparticles (0.33 mmol Fe3O4), 0.83 g (2.2 mmol) of Au(OOCCH3), 3.1 g (12 mmol) of 1,2-hexadecanediol, 0.5 mL (∼1.5 mmol) of oleic acid, 3 mL (∼6 mmol) of oleylamine were added into 30 mL of phenyl ether. In this case, the mole ratio of the Au precursor to the iron oxide nanoparticles was approximately 7:1. Under a nitrogen atmosphere the reaction solution was heated to 180−190 °C and was kept at this temperature for 1.5 h. After cooling to room temperature, ethanol was added into the solution. A dark purple material was precipitated and separated by centrifuging. The precipitated product was washed with ethanol and dispersed in hexane. After the synthesis, we transferred the organic NPs in water by addition of methoxy-PEG thiol/ss15mer thiol with a 1:1 ratio. We performed the phase transfer by layering 100 μL of NP dispersion on a solution (1 mL, 5 mM) of methoxy-PEG thiol/ss15mer and keeping it under agitation for 24 h. Then, the organic phase was removed by evaporation of the solvent, and the NPs in the water phase were washed twice by centrifuging at 10000g/30 min to remove the unbound ss-DNA. Then, a cholesteryl DNA (chol-DNA) with a sequence half complementary to a portion of the zipper was added. The NPs with the ss-DNA, the chol-DNA, and the complementary sequence were mixed at room temperature for 24 h. Then the dispersion was annealed from 15 °C to 60 °C at 1 °C/min. The platform as prepared was then conjugated to liposomes by addition of 25 μL (0.9 nmol) of NPs-DNA dispersion to 200 μL of a liposome dispersion (0.25 mg/mL) (reaching a more or less 30 NPs-DNA/ liposome ratio). The same procedure was adopted for magnetoliposomes. Preparation of Liposomes and Magnetoliposomes. For the preparation of lipid carriers, first a DPPC and DPPC-NP dry film was obtained, by evaporation of the solvent from a CHCl3 solution of the lipid (DPPC) and an aliquot of Fe3O4 NPs coated with oleic acid in hexane (15 μL of 8 mg/mL NP dispersion); the dry film was hydrated with water to a final lipid concentration of 1 mg/mL. Regarding bare vesicles, the suspension was homogenized with vortex mixing until complete dispersion of the film and frozen−thawed 10 times; then multilamellar polydisperse vesicles were extruded at 60 °C through a polycarbonate membrane of 0.1 μm pore size (Whatman). On the other hand, an ML lipid film was hydrated with water, frozen−thawed 10 times, and then tip sonicated. For the release study, 10 mM carboxyfluorescein was used as a small model drug, instead of deionized water. Nonencapsulated CF was removed by dialyzing the sample for 3 days. Lipid and nanoparticle concentrations were measured by inductively coupled plasma−atomic emission spectrom-

g2(q , τ ) = 1 + β |g1(q , τ )|2

(1)

with β being the spatial coherence factor, which depends on the geometry of the detection system. The field autocorrelation functions were analyzed through a cumulant analysis stopped to the second order. Measurements have been performed at 25 °C on diluted sample to avoid multiple scattering. Fluorescence Correlation Spectroscopy. FCS experiments were carried out with a Leica TCS SP2 laser scanning confocal microscope (Leica Microsystems GmbH, Wetzlar, Germany) equipped with a 63× water immersion objective instrument by using the 488 nm laser line. The fluorescence emission was acquired with an ISS module (ISS, Inc., Champaign, IL, USA) equipped with an APD with 500−530 nm (where the fluorescence emission of Oligreen was acquired) band-pass. The autocorrelation function of the fluorescence intensity, G(τ), is evaluated as a function of the fluctuations of the signal from the average value, as

G(τ ) =

δI(t ) δI(t + τ ) ⟨I(t )⟩2

(2)

The models employed for the analysis of the autocorrelation functions take into account the shape and the exact size of the detection volume, which is approximated as a 3D-ellipsoidal Gaussian shape with axial (z0) and lateral (w0) defining parameters, determined through a calibration performed with a reference fluorescent dye with known diffusion coefficient (50 nM standard solutions of Rhodamine 7757


Article

ACS Nano 110 with D = 430 μm2 s−1 in water at 25 °C). For a three-dimensional Brownian diffusion mode in a 3D Gaussian volume shape, the ACF profiles can be analyzed according to the equation64,65

G(τ ) =

−1 −1/2 ⎛ 4Dτ ⎞ ⎛ 4Dτ ⎞ ⎜ ⎟ ⎜ ⎟ + + 1 1 w0 2 ⎠ ⎝ z02 ⎠ ⟨c⟩π 3/2w0z 0 ⎝

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1

(3)

with ⟨c⟩ the average fluorophore concentration and D the diffusion coefficient of the probe (μm2 s−1). The general expression for the analysis of the autocorrelation function of the fluorescence intensity due to the 3D diffusion contribution of different i diffusing entities is

G(τ ) =

1 ⟨c⟩π 3/2w0z 0

⎛

∑ fi ⎜1 + i

⎝

−1 −1/2 4Diτ ⎞ ⎛ 4Diτ ⎞ ⎟ ⎜ ⎟ + 1 w0 2 ⎠ ⎝ z02 ⎠

(4)

with ⟨c⟩ being the average fluorophore concentration and f i the weight factor of each i diffusing component characterized by diffusion coefficient Di. For FCS measurements on the ON and its insertion inside liposomes, 2 μL of Oligreen was added to 125 μg/mL sample dispersions. Fluorescence Leakage Studies. Steady-state fluorescence was measured with an LS50B spectrofluorimeter (PerkinElmer, Italy). The emission fluorescence spectra of CF were recorded between 500 and 610 nm with excitation wavelength set at 492 nm. Fluorescence experiments were also performed on both magnetoliposomes and liposomes at a concentration of 150 μg/mL before and after AMF exposure, at regular time intervals. All samples were diluted to the measurement concentration with a solution of Triton X-100 1% v/v to achieve complete release of CF through vesicle disruption. The release percentage was calculated from the fluorescence intensity as

%release =

[F(t ) − F0] × 100 [Ffinal − F0]

(5)

where F(t) is the intensity as a function of the time, F0 is the initial intensity, and Ffinal is the intensity after the addition of Triton X-100 1%.

ASSOCIATED CONTENT S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsnano.6b03194. AMF generator specifications; fitting models for SAXS analysis; FCS extended data (PDF)

AUTHOR INFORMATION Corresponding Author

*Phone: +39-055-4573038. E-mail: debora.berti@unifi.it, [email protected]fi.it. Notes

The authors declare no competing financial interest.

ACKNOWLEDGMENTS All the authors acknowledge financial support from CSGI and MIUR through the project PRIN 2010- 2011 grant 2010BJ23MN. All the authors would like to thank Mirko Severi for ICP measurements and Lorenzo Mannelli for designing the ERB B2 antisense oligonucleotide. REFERENCES (1) Sapsford, K. E.; Algar, W.; Berti, L.; Gemmini, K.; Casey, B.; Oh, E.; Stewart, M.; Medintz, I. Functionalizing Nanoparticles with Biological Molecules: Developing Chemistries that Facilitate Nanotechnology. Chem. Rev. 2013, 113, 1904−2074. 7758


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ACS Nano

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