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Tunable hydrogel—Nanoparticles release system for sustained combination therapies in the spinal cord
Colloids and Surfaces B: Biointerfaces 108 (2013) 169–177
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Colloids and Surfaces B: Biointerfaces journal homepage: www.elsevier.com/locate/colsurfb
Tunable hydrogel—Nanoparticles release system for sustained combination therapies in the spinal cord Filippo Rossi a,b , Raffaele Ferrari a , Simonetta Papa b , Davide Moscatelli a , Tommaso Casalini a , Gianluigi Forloni b , Giuseppe Perale a,b,∗ , Pietro Veglianese b a b
Politecnico di Milano, Dipartimento di Chimica, Materiali e Ingegneria Chimica “Giulio Natta”, via Mancinelli 7, 20131 Milano, Italy Istituto di Ricerche Farmacologiche Mario Negri-IRCCS, Dipartimento di Neuroscienze, via La Masa 19, 20156 Milano, Italy
a r t i c l e
i n f o
Article history: Received 28 December 2012 Received in revised form 26 February 2013 Accepted 27 February 2013 Available online 14 March 2013 Keywords: Drug delivery Hydrogel Polymeric nanoparticles Release system Spinal cord
a b s t r a c t Poly(methyl methacrylate) (PMMA) nanoparticles (NPs) were prepared by emulsion free radical polymerization. NPs with controlled dimension, as monitored by dynamic light scattering (DLS) and transmission electron microscopy (TEM), were produced by changing experimental parameters, such as the amount of emulsifier and the monomer feeding mode (batch or semi-batch). Then, different sized NPs (60, 80 and 130 nm) were loaded in polysaccharide-polyacrylic acid based hydrogels, cross-linked by covalent ester bonds between polyacrylic acid (PAA) and agarose chains, with different pore sizes (30, 60, 90 nm). The characteristics of the resulting composite hydrogel-NPs system were firstly studied in terms of rheological properties and ability to release Rhodamine B that presents steric hindrance similar to many neuroprotective agents used in spinal cord injury (SCI) repair. Then, diffusion-controlled release of different sized NPs from different entangled hydrogels was investigated in vitro and correlated with NPs diameter and hydrogel mean mesh size, showing different hindrances to the diffusion pathways. Release experiments and diffusion studies, rationalized by mathematical modeling and verified in vivo, allowed to build a material library able to satisfy different medical drug delivery needs. © 2013 Elsevier B.V. All rights reserved.
1. Introduction Traumatic spinal cord injury (SCI) is a dramatic condition that results in profound and life-lasting disability for the patient, with consequent reduction of the quality of life. Its incidence in the EU ranges between 28 and 55 cases per million people, with about 11.000 new cases reported every year [1–3] in the sole Europe. The lesion mainly results from a contusive, compressive or stretch injury of the cord. However, most of the post-traumatic degeneration of the tissue is due to the so-called secondary injury, which is known to be a multifactorial process [4–7]. In the last decades the primary interest of the scientific community has been pointing toward the mechanisms of secondary injury in order to minimize its pathological consequences [8,9]. Indeed, a growing number of pharmacological treatments have been proposed as potential therapeutic approaches, but nowadays none of them showed evident therapeutic efficacy when translated to humans [10]. Probably this is due to the fact that several therapeutic approaches aim to counteract only one single effect, while different therapeutic targets
∗ Corresponding author at: Dipartimento di Chimica, Materiali e Ingegneria Chimica “Giulio Natta”, Politecnico di Milano, via Mancinelli 7, 20131 Milano, Italy. Tel.: +39 02 2399 3145; fax: +39 02 2399 3180. E-mail address: [email protected] (G. Perale). 0927-7765/$ – see front matter © 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.colsurfb.2013.02.046
should be considered at the same time: multifunctional therapies directed to counteract multiple injury mechanisms, trying to combine both neuroprotective and neuroregenerative agents, seem to be extremely promising [3,4,11,12]. Hence, the ability to guarantee different controlled drug delivery kinetics using neuroprotective agents in the first days and neuroregenerative in the following ones could be an efficacious therapeutic strategy [13–15]. In this framework, recent advances in polymer science have provided a huge amount of innovations, underlining the increasing importance of macromolecules in controlled release applications [16–20]. Particularly NPs, thanks to their versatility in terms of size, potential surface and hydrophilic/lipophilic characteristics, lead relevant advantages in drug delivery by increasing the selectivity of drugs and by controlling their release during time [21–25]. Hence, a good strategy could be to load drugs within NPs in order to deliver hydrophobic drugs into injured spinal cord for therapeutic purposes [8,26]. Direct injection of colloidal NPs suspension is one attractive alternative, but several concerns arise: injected NPs very often leave the zone of injection as they are not confined by any support, and easily extravasate into the circulatory torrent, migrating all over the body to the liver and the spleen, or toward an uncertain faith [8,27]. According to these critical issues, nanostructured hydrogels could be reliable NPs carriers, due to their ability to control release rates in situ [13,28,29].
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In this framework, we investigated the potential use of PMMA NPs loaded within hydrogel networks, hence forming composite materials [1], for drug delivery purposes in SCI repair strategies. In particular it should be underlined that NPs directly injected within the spinal cord are quickly (i.e. few hours) uptaken by macrophages [30,31]. Indeed, the controlled release of drug loaded NPs from hydrogels could help to both sustain a selective and effective release of drugs and also delay the cell uptake mechanism. The chosen hydrogel library, specifically developed for SCI repair strategies, was obtained by synthesis from statistical block polycondensation between agarose and carbomer 974P, together with cross-linkers (being a library of hydrogels, they differ by means of components mutual ratios and hereafter they are briefly termed as AC, acronymic of their main components) [27,32]. In previous works we observed that our hydrogels can remain localized at the site of injection [27], showing high biocompatibility and good ability to provide short term delivery for neuroprotective molecules together with long term for neuroregenerative ones [14]. PMMA NPs were synthesized through emulsion free radical polymerization process, which ensures a very narrow particle size distribution and an easy control over final average particle size [33,34]. Then, after AC synthesis but before sol/gel transition takes place, NPs were physically entrapped within the AC networks forming ACNPs composite materials. Since PMMA based NPs are not easy biodegradable, they can be traced both in in vitro and in vivo biological studies over a long period of time without side effects due to the biodegradation process. Obtained composite materials were characterized from a physical-chemical point of view and compared with neat AC hydrogels, already studied in previous works [32]. NPs release from AC hydrogels was studied with the aim to develop a new library of hydrogel-nanoparticles based materials able to provide different release rates, tuning independently both hydrogel mean mesh size (30, 60, 90 nm) and NPs diameter (60, 80, 130 nm): ability to provide different release kinetics was demonstrated and also verified in vivo in a spinal cord mouse model. 2. Experimental 2.1. Materials For NPs synthesis, methyl methacrylate (MMA, 99% purity, Sigma–Aldrich, Germany), polyoxyethylene sorbitan monooleate (Tween80, Sigma–Aldrich, Germany) and potassium persulfate (KPS, >99% purity, Sigma–Aldrich, Germany) were all used as received. For hydrogel synthesis: cross-linked polyacrylic acid carbomer 974P was purchased by Fagron (The Netherlands), phosphate buffered saline solution (PBS), propylene glycol, glycerol and sodium hydroxide were purchased by Sigma–Aldrich (Germany) while agarose from Invitrogen (Carlsbad, CA, USA) and all were used as received. Rhodamine B base (RhB) (Sb sensitivity <0.1 g mL−1 , Carlo Erba reagents, Italy) was used for release experiments. 2.2. NPs synthesis Emulsion free radical polymerization reaction of MMA was carried out in batch and semi-batch conditions, as already described [35], using a 100 mL three necked glass flask equipped with a reflux condenser; temperature were controlled with an external oil bath. Different amounts of Tween80 were added to 100 mL of deionized water and the solution was heated up to 80 ◦ C. Vacuum-nitrogen cycles have been used to make and to maintain inert the reaction atmosphere. The initiator, KPS, in concentration of 1% wtKPS /wtMMA was added to the purged solution. Next, 5 g of MMA was injected with a flow rate of 3 mL/h using a syringe pump (Model NE-300, New Era Pump System, Farmingdale, NY, US) for semi-batch
reaction, while for batch reaction 5 g of MMA were all injected before the initiator. In both cases reaction was run for 3 h. Final monomer conversion was measured through gravimetric measurement and calculated as higher than 98.5%. Different particle sizes were obtained varying the emulsifier-monomer ratio and the feeding condition (batch and semi-batch). In particular, NPs of 60 and 80 nm were produced using semibatch conditions with 30% and 20% wtTween80 /wtMMA ratio, while NPs of 130 nm were obtained with a 15% wtTween80 /wtMMA ratio in batch conditions. RhB loaded NPs of 130 nm were produced using the same experimental condition; where MMA and RhB (5 mg) solutions were fed into the reactor. RhB concentration was measured by spectrofluorimetry at excitation (ex) and emission (em) wavelengths of 552 and 585 nm, respectively, using a Tecan Safire plate reader (Durham, NC). Percent encapsulation efficiency (% EE) was calculated based on the following equation: % EE =
Drug entrapped in NPs · 100 Initial amount of drug added
(1)
2.3. Characterization of NPs Final particle size has been determined by DLS (Malvern, Zetanano ZS, US); analyses were performed in triplicate and reported data are the average of three runs where the standard deviation was always below 5%. Dimension and morphology of produced NPs were also confirmed by transmission electron microscopy (TEM; using a EFTEM Leo 912AB, at 80 kV, by Karl Zeiss, Jena, Germany). Samples were prepared placing 5 L drop of NPs dispersion on a Formvar/carbon-coated copper grid and dried overnight. Digital images were acquired by a charge coupled device (CCD; Esi Vision Proscan camera). 2.4. Hydrogel synthesis Polymeric solution was achieved by mixing polymer powders of agarose and carbomer 974P in PBS (Phosphate Buffer Saline solution) in batch condition, adding a mixture of cross-linking agents made of propylene glycol and glycerol along with NaOH 1 N (reaction pH was kept neutral) [14,32]. The gelation onset was achieved by means of electromagnetic stimulation (500 W irradiated power) heating in ratio of 1 min per 10 ml of polymeric solution at 80 ◦ C. Microwave-enhanced chemistry is based on the efficient heating of materials by “microwave dielectric heating” effects [36]. This phenomenon is dependent on the ability of a specific material (solvent or reagent) to absorb microwave energy and convert it into heat. This efficient method could allow gel forming in less than 10 min, avoiding prolonged heating that could damage the structures of the components and the compounds loaded. Here, carbomer carboxyl groups constitute the main cross-linking sites to be reacted with hydroxyl groups from agarose, altogether giving rise to the three dimensional matrix. Propylene glycol and glycerol take part in the condensation reaction and form local interconnections through their own hydroxyl groups [14]. They also play a physical role, increasing system viscosity and thus enhancing tight macromers cross-linking. AC formulations are numbered according with previous works [35] and their detailed compositions are presented in Supplementary Material. The differences in these three formulations are due to the different presence of cross-linkers, resulting in different mesh size values: nominal 90 nm for AC1, 60 nm for AC3 and 30 nm for AC6 [37]. By varying the amounts and the functionality, f in Flory notation, of the cross-linkers we can synthesize more or less entangled polymeric networks with high impact on physico-chemical properties [37]. Thus, acting on microchemistry of hydrogel library and following a multi scale approach, it was possible to manage macro properties.
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2.5. Rheology Rheological analyses were performed, on AC-NPs systems, at 37 ◦ C, using a Rheometric Scientific ARES (TA Instruments, New Castle, DE, USA) equipped with parallel plates of 30 mm of diameter and a 4 mm gap between. Oscillatory responses, dynamic frequency sweep tests (Gˇı, elastic modulus, and Gˇıˇı, loss/viscous modulus) were determined at low strain values over the frequency range 0.1–500 rad/s [37]. Thixotropic behavior of the samples was also investigated and shear strain and viscosity, as a function of shear rate, were evaluated too [32]. To evaluate the role of hydrogel we performed rheological studies with AC1, AC3 and AC6 loaded with NPs (80 nm), here briefly named AC1-NPs80, AC3-NPs80 and AC6NPs80. The role of NPs diameter was studied comparing AC1-NPs80 with AC1-NPs130. 2.6. Rhodamine B release from AC hydrogel and AC-NPs composite system RhB was chosen mainly because of its steric hindrance, similar to many small drugs, and for its absorbance that makes it easily detectable by UV spectroscopy [11]. It was loaded within (i) AC1 hydrogel network and (ii) within NPs (130 nm) then loaded in AC1 hydrogels (AC1-NPs130) to compare the release ability of these two systems. (i) RhB aqueous solution (0.1 mg/mL) was mixed with AC1 gelling solution, above sol/gel transition to allow good solute dispersion within polymeric network; gelation took place in steel cylinders (0.5 cm3 , d = 1.1 cm). (ii) NPs (loaded with RhB as explained in Section 2.2) solution was mixed with AC1 gelling solution, above sol/gel transition to allow good solute dispersion within polymeric network. In order to allow good comparison between these two systems, RhB concentration was kept constant in both of them. Gelation took place in steel cylinders (0.5 cm3 , d = 1.1 cm). Then, three samples for each condition were put in excess of cerebrospinal fluid (CSF) and aliquots were collected at defined time points, while the sample volume replaced by fresh CSF, in order to avoid mass-transfer equilibrium between the gel and the surrounding solution. Percentage of RhB released was measured by UV spectroscopy [38]. 2.7. NPs Release from AC hydrogels and mathematical modeling AC1, AC3 and AC6 hydrogels were loaded with different PMMA NPs, in terms of diameter (60, 80, 130 nm), to assess their in vitro diffusion-controlled delivery. Composite materials were briefly named as following: AC1-NPs60 for AC1 hydrogel loaded with 60 nm diameter NPs etc.. . . During AC hydrogels cooling phase, slightly above 37 ◦ C, AC samples were mixed with the chosen NPs solutions at 1% monomer concentration: this procedure allows solute loading still during sol state, i.e. before sol/gel transition [14]. The loaded AC-NPs systems were cast in standard plastic 48well cell culture plates, 0.5 mL/well (diameter = 1.1 cm). Complete swelling was achieved leaving samples in a NPs isotonic solution within a standard incubator at 37 ◦ C and 5% CO2 atmosphere, overnight. This procedure is necessary to allow swelling without having any NPs release. Swelling equilibrium was quickly reached in the first hour and, on the other hand, stability of this material was quite high: it remained stable for weeks before significant degradation, by hydrolysis, took place [37]. Hence, as characteristic time is much larger than experimental length, degradation reactions can be taken as fully neglectable in this framework. Therefore, NPs flux is controlled only by concentration gradient and diffusion can be considered as Fickian. Hydrogels were then placed in CSF filled wells (5 mL volume each). Aliquots were collected at defined time points, while the sample volume replaced by fresh CSF, in order to avoid mass-transfer equilibrium between the gel and the
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surrounding solution. Samples collected were analyzed to assess the NPs release percentage using DLS and full details are contained in Supplementary Material. In addition, to fully recapitulate this AC-NPs composite material library, NPs diffusivity (D) was evaluated with a mathematical model based on mass balances, i.e. on fundamental conservation laws [32,39]. Diffusion is described through the second Fick law with a 1-dimensional model in a cylindrical geometry, as shown in Eq. (2). This particular hydrogel shape was chosen to best fit the average cystic cavity present inside an injured spinal cord [1,39]. Here radius (r) is the characteristic dimension for the investigated transport phenomenon. Therefore, the above-mentioned increase takes place due to the material flux, which takes place at the CSF/hydrogel surface. Eqs (6) and (7) represent the boundary conditions for the left and the right border, respectively. The first one implies profile symmetry at the center (that is, with respect to cylinder axis), while the second one represents the equivalence between the material diffusive fluxes at the CSF/hydrogel surface. ∂CG 1 ∂ =D· · · r ∂r ∂t VS ·
r·
∂CG ∂r
∂CS = kC · Sexc · (CG − CS ) ∂t
CS (t = 0) = 0
−D ·
mG,0 VG
=0 r=0
(3) (4)
CG (t = 0) = CG,0 = ∂CG ∂r
(2)
(5)
(6)
∂CG ∂r
= kC · (CG − CS )
(7)
r=R
The two mass balance equations involve the mean NPs concentration within the hydrogel (CG ), the mean NPs concentration in the outer solution (CS ), the volume of the solution (VS ), the hydrogel volume (VG ), the NPs mass present inside the matrix (mG ) and the exchange interfacial surface (Sexc ), i.e. the boundary surface between gel and surrounding solution (which, simplifying, can be here considered as being only the side surface). Finally, D represents the diffusion coefficient and kC , the mass transfer coefficient. The mass transfer coefficient is computed through Sherwood number obtained by means of penetration theory [40]: Sh =
8 kC · 2r = D
(8)
2.8. Animals and their care C57BL/6 mice were used for this study. Procedures involving animals and their care were conducted in accordance with the institutional guidelines, which are in compliance with Italian national (D.L. no. 116, G.U. suppl. 40, Feb. 18, 1992, Circolare No.8, G.U., 14 luglio 1994) along with international laws and policies (EEC Council Directive 86/609, OJL 358, 1 DEC.12, 1987; NIH guide for the Care and use of Laboratory Animals, U.S. National Research Council, 1996). All animals were housed under standard conditions (22 ± 1 ◦ C, 60% relative humidity, 12 h light/dark schedule), one per cage, with free access to food (Altromin MT, Rieper, Vandoies (BZ), Italy) and water. 2.9. Surgery To obtain a reliable translational study and to investigate different NPs release possibilities in vivo, twelve mice were enroled for the surgery and randomly divided into 2 groups of 6 animals each:
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Table 1 Characteristics of PMMA based NPs. Nominal diameter [nm]
Diameter [nm]
PSD
-potential [mV]
60 80 130
63.80 81.52 131.9
0.063 0.062 0.070
−27.1 −22.8 −23.3
animals received AC1-NPs60 and AC1-NPs130 respectively. The NPs concentration loaded is the same used for in vitro studies. All surgery procedures were performed under deep anesthesia by 2.5% Avertin (2,2,2-tribromoethanol in 2-methyl-2-butanol, 400 mg/kg, i.p., Sigma–Aldrich, Germany). Thirty minutes before undergoing surgery, animals received an injection of ampicillin (0.03 mg/kg, subcutaneous, Roche, Italy) and buprenorphine (0.05 mg/kg, subcutaneous, Sigma–Aldrich, Germany). The back of the animal was shaved at dorsal level and a cutaneous incision (3 cm) was performed to expose the backbone. T11 and T12 vertebrae were identified and exposed by separation of dorsal and intervertebral muscles. Animals were then placed in a Cunningham Spinal Cord Adaptor (Stoelting, Ireland) mounted on a stereotaxic frame, and laminectomy of the T11 vertebra was done to uncover thoracic spinal cord. Subsequently, a gelling solution (6 L) was placed with a micropipette on top of the cord in the subdural space [14]. After gel positioning, dorsal muscles were juxtaposed using absorbable sutures and the skin sutured and disinfected. After surgery animals were kept on a warm pad for 30 min and then placed in separated cages for recovery. After seven days animals were anesthetized with Equithesin (1% Phenobarbital, 4% chloral hydrate, 30 mL/10 g, i.p.) and transcardially perfused with 20 mL saline and subsequently with 50 mL of sodium phosphate-buffered 4% formaldehyde solution. Spinal cords were rapidly removed and the hydrogels recovered and analyzed with DLS to evaluate the percentage of NPs present within them. 2.10. Statistical analysis Where applicable, experimental data were analyzed using Analysis of Variance (ANOVA). Statistical significance was set to p value <0.05. Results are presented as mean value ± standard deviation. 3. Results and discussion 3.1. Nanoparticles synthesis As far as emulsion free radical polymerization is a wellestablished process, it is possible to tune the final NPs characteristics by changing process parameters such as the amount of emulsifier, the reaction temperature and the feeding mode [41,42]. It is worth noticing that emulsion polymerization ensures the final biocompatibility of the NPs suspensions: indeed, the reaction is carried out in water without using any solvent and adopting a FDA approved emulsifier such as Tween80 [34]. Beside biocompatibility, this emulsifier was chosen, despite not being as efficient as ionic surfactants (like e.g. sodium dodecyl sulfate), because it does not introduce any charge effects that might affect the NPs release from the hydrogel matrix. Finally, this technique allows the synthesis of polymeric NPs with narrow particle size distribution (PSD), necessary to discriminate NPs with relatively similar particle size. Obtained results are shown in Table 1, where NPs were produced with the process parameter reported in the Experimental Section. As expected, NPs obtained with different amount of emulsifier and feeding mode are different in size. In particular, batch operation mode leads to larger NPs, namely 130 nm, while smaller NPs are produced adopting semi-batch process. Accordingly,
maintaining semi-batch condition, a lower amount of emulsifier leads to bigger NPs. Moreover, PSD is very narrow, with values lower than 0.1 (Table 1), as observable from the DLS and TEM images of the NPs, presented in Supplementary Material, which shows how NPs present similar sizes; some aggregates are detected but they are due to the sample preparation technique which leads to aggregates formation, anyway each single particle is very well visible. As a final remark, -potential data are in the expected range of NPs produced using steric emulsifier as Tween80, where the charge of the NPs come only from the initiator, KPS [43].
3.2. Rheology The storage modulus, G , was compared between unloaded gel samples (AC1) and NPs loaded ones (AC1-NPs80) to investigate the effect of NPs on the rheological behavior of these composites (Fig. 1A). In previous works gel storage modulus (G ) was found to be approximately one order of magnitude higher than the loss modulus (G ), indicating an elastic rather than viscous material [37]. Furthermore, both G and G of each gel were essentially independent from frequency over the entire investigated range, thus indicating dominant viscoelastic relaxations of networks at lower frequencies. Such rheological behavior matches the characteristic signature of a solid-like gel, confirming this nature for AC type gels (see Supplementary Material). These values are in complete accordance with literature and, in particular, with the requirements for SCI repair applications [1]. The presence of NPs does not affect the rheological properties of the composite system and also in this case G trends are essentially independent from frequency over the range investigated. In comparing neat systems with those containing NPs, G increase is evident: specifically, the addition of NPs immediately results in a 3-fold increase in G over neat AC gels, thus suggesting a stiffer and more elastic material. The increased elastic modulus could be attributed to interactions between AC network and NPs. In particular electrostatic interactions between negatively charged AC network and negatively charged PMMA NPs increase the chain mobility, hence increasing the whole system elastic properties. The consistent and monotonic enhancement in hydrogel mechanical properties also suggests that NPs are well dispersed within the polymeric matrix. In addition, AC hydrogels do not show any loss in elasticity upon addition of NPs, suggesting that the underlying polymer network remains intact upon NPs addition and that it does not change the morphology and the shape of loaded NPs. This assumption is further corroborated by evidences from DLS (see Supplementary Material, Table S.2). Following the oscillation frequency sweep, the thixotropic loop method with a flow shear-rate sweep was performed on both neat and NPs-loaded AC gels results are presented in Fig. 1B: the typical area visible from thixotropic loop method corresponds to the energy needed to breakdown the structural stability of the material and the time necessary for the rearrangement of the material after being sheared. Once the structural effect on rheological properties was confirmed, the NPs loading contribution was considered: hysteresis loop areas are bigger for both NPs-loaded hydrogels with respect to neat ones. The presence of NPs, indeed, makes the gels more stable and leads to higher mechanical properties, in agreement with dynamic frequency sweep tests. In summary, NPs loaded composite materials, compared with neat ones, show faster sol/gel transition and this effect could be a relevant improvement for injectability [44]. Similar results were obtained also with AC3 and AC6 gels, here not presented for bare sake images clarity (see Supplementary Material, Fig. S.3). The same investigations were also performed considering NPs of all different sizes: as predictable, gel elastic modulus G shows a high dependence on NPs diameter, increasing with NPs size increase (see Supplementary Material).
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Fig. 1. Rheological analyses: (A) dynamic frequency sweep test of composite material at 37 ◦ C with small oscillatory shear in the linear viscoelastic regime: (A) G ( AC1, 䊉 AC1-NPs). (B) Thixotropic loop for: AC1 () and AC1-NPs (䊉).
3.3. Rhodamine B release Once proved that the AC-NPs composite systems maintain adequate mechanical/structural properties, RhB release was investigated from both neat AC and AC-NPs hydrogels. Firstly, the in vitro drug release kinetics of RhB, loaded in neat (a) AC1 hydrogel and in (b) NPs (130 nm) then loaded in AC1 hydrogels (AC1-NPs130), were deeply investigated and compared. The importance of this study resides in the necessity to prove NPs ability to release drugs with proper kinetic on one side and, on the other, to verify the advantages behind the combination of hydrogels together with NPs in a composite release system. As said, RhB was chosen for this investigation because its steric hindrance resembles many corticosteroids and anti-inflammatory drugs (for example methylprednisolone and estradiol), commonly used in pharmacotherapy to hinder SCI secondary effects and CNS damages [8]. RhB release profiles from NPs and AC1-NPs130 systems are compared in Fig. 2: from Fig. 2A it is well visible that RhB release profile from AC1 samples is very fast and it is almost completed in the first 12 h, while RhB release from AC1-NPs130 system is prolonged for a much longer period; from Fig. 2B it is also evident that loading NPs within hydrogels is hence the most promising solution. There the influence of the system in delivering RhB was investigated plotting release percentage against time square root. A linear plot is indicative of Fickian diffusion and the y-axis intercept value an indication of
burst release, where it is well known that an ideal controlled release system should present linear trend during time and its y-axis intercept equal to zero [45]. RhB directly loaded within AC hydrogels shows a linear trend only in the first hours and then a plateau trend is visible. Moreover, in this case the burst release contribution is high (about 40%), underlining the expected poor ability to control release of small drugs [14]. On the other hand, loading NPs, that contain RhB, within AC hydrogels increases performances in terms of pure diffusive mass transport and of burst release contribution that results to be extremely small (around 1%). This is due to the fact that AC and NPs represent multiple diffusion barriers to RhB release and so the AC1-NPs130 system can control and sustain better the release of small drugs. Thus, after having verified the increased release ability of small drug from AC1-NPs system, we investigated the possibility to build an AC-NPs material library that can tailor the diffusion during time of NPs and, consequently, of the loaded drugs. 3.4. NPs release from hydrogel We discussed above how a small drug mimetic (RhB) can be released from AC-NPs composite systems and hence the advantages to load NPs within AC hydrogels. We further investigated the possibility to build a composite material library loading NPs with different diameter within hydrogels with different pore size,
Fig. 2. (A) In vitro release profiles of RhB from AC1 hydrogel () and AC1-NPs130 composite system (䊉) in CSF at 37 ◦ C. (B) The slope of RhB releases against the square root time is representing how the delivery system influences Fickian diffusion.
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Fig. 3. (A) Cumulative release of NPs with different diameters (red = 130 nm; blue = 80 nm; black = 60 nm) from AC1 hydrogel. The same trends were also found with AC3 and AC6 formulations here not presented for image clarity; (B) release of NPs (diameter = 80 nm) from AC1 () and AC6 () hydrogels. (C) The slope of NPs (red = 130 nm; blue = 80 nm; black = 60 nm) release from AC1 against the square root of time is representing how NPs diameter influences Fickian diffusion. (D) The slope of NPs release from AC1 (), AC3 () and AC6 () against the square root of time is representing the role of hydrogel mesh size in mass transport. Data points are presented mean ± standard deviation of three different gel cylinders. (E) Effect of NPs diameter and AC hydrogel mean mesh size on NPs diffusion coefficient: D trend exhibits high dependence on NPs diameter (red = 130 nm; blue = 80 nm; black = 60 nm) while AC mean mesh size is less able to influence mass transport through the network pores. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of the article.)
evaluating their release during time. Our previous works [14,32] have demonstrated that solutes physically entrapped within the network leave completely hydrogel matrix with kinetics depending on their hydrodynamic radius: mass flux is, indeed, driven only by concentration gradient and diffusion can be considered as Fickian, with no loss of generality as long as gel degradation characteristic time is much higher than diffusion one. Here, to investigate the influence of particle size on the release kinetics, different sized NPs were loaded within different hydrogel networks at sol state, just before sol/gel transition. In Fig. 3A the release kinetics from AC1 gel formulation with three different NPs diameters are
presented. The same trends were also found with AC3 and AC6 formulations here not presented for bare sake of image clarity (see Supplementary Material, Fig. S.5). The percentage of NPs released from AC-NPs in the first 2 h, less than 30%, is attributed to the high initial concentration gradient and in particular at NPs that were at or near the solvent–hydrogel interface and thus could rapidly diffuse into the supernatant solution. Thereon, NPs were slowly released showing a diameter dependent kinetic: the bigger NPs need more time to be released from the hydrogels than the smaller ones. In Fig. 3B, on the other hand, the release kinetics of the same NPs (80 nm) from different AC formulation (AC1 and AC6) are
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Fig. 4. (A) The AC gelling solution was homogenized with NPs solution (red 130 nm, yellow 60 nm) above the sol–gel transition temperature and then injected into the spinal cord. After 7 days AC-NPs was recovered and analyzed with DLS to quantify the in vivo percentage of NPs with different diameters (60 nm and 130 nm) released from AC1 hydrogel; (B) picture of in situ gelified AC-NPs composite system after 7 days before hydrogel recovering (white rectangle; scale bar = 0.5 mm); (C); percentages of NPs60 and NPs130 released from AC1 hydrogels after 7 days. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of the article.)
reported. AC1 hydrogel generally showed the fastest release when compared to all other formulations. From Fig. 3A and B it is evident that hydrogel mesh sizes is not the limiting factor for drug release while NPs diameter is the real limiting parameter. This is due to the fact that the calculated hydrogel mesh size is an average value [28]: they are synthesized from statistical random walk reaction and so they are not polymeric ordered matrices. Moreover, to better investigate and understand the influence of NPs diameter and AC mesh sizes, release percentage is plotted against time square root (see Fig. 3C and D). NPs 60 and 80 nm showed linear behavior only in the first hours (Fig. 3C), while for NPs 130 nm this is evident for the whole considered timeframe. This difference is due to the initial contribution of burst release that is small for low NPs values and is not present considering NPs 130 nm. Hence, increasing NPs diameter, initial burst release could be limited and a linear trend for all experimental times could be provided. On the opposite, as explained above, AC mesh size role in NPs delivery tuning is less important: from Fig. 4D is visible that, changing mean mesh size, the mass transport is only partially influenced. So, the investigation of parameters roles allows high drug release kinetic tunability according to different potential therapeutic approaches in SCI repair. In particular it should be underlined that the most probable reason behind disappointing results in SCI treatment is due to the fact that SCI is characterized by a temporal development of biochemical pathways of degeneration. Recent research has focused its attention on multifactorial therapies
directed to counteract multiple injury mechanisms, trying to combine both neuroprotective and neuroregenerative agents [12,46]. In particular, as explained above, controlling the release of NPs-drug loaded consents a selective and sustained treatment. 3.5. NPs diffusivity evaluation Mass release data obtained experimentally (Fig. 3A–D) were used to estimate NPs diffusion coefficients. As explained above, the release mechanism could be considered as a pure Fickian diffusion, being concentration driven through AC hydrogel pores and allowing to neglect any contribution of both swelling and degradation phenomena. Fig. 3E shows the dependence of NPs diffusivity upon their diameter and AC mean mesh size. The diffusion coefficient, magnitude order of 10−6 cm2 /s, does not show any high differences among the three AC formulations studied, even if materials are structurally and mechanically different. In particular, it should be noted that NPs with average diameter of 130 and 80 nm have bigger hindrances with respect to the mean AC mesh sizes and so their diffusivity trends are similar: no evident differences can be observed if we consider an iso-NPs diameter line in the graph. On the other hand, if we consider 60 nm NPs, a slope increase is well visible between mesh size values of 30 and 60 nm: this is reasonable because decreasing the entanglement of the matrix (from AC6 to AC3) the diffusion is favored. In general it seems that, as described above, NPs mobility through the pores is affected only
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partially by the surrounding hydrogel inner network conformation and, hence, mean mesh size represents only a partial hindrance to NPs diffusion pathway. In contrast, NPs diffusivity showed high dependence on particle size with a decreasing trend from small diameters to big ones. In summary, NPs diffusivity values are in accordance with expectations deducible from previous works [32], exhibiting lower values respect to the drug directly loaded in neat AC hydrogels, due to their higher hydrodynamic radius. This result is, indeed, a direct consequence of the Stokes-Einstein equation, which describes diffusivity in inverse proportion to hydrodynamic radius of released molecule. Nevertheless, it is evident that the two parameters involved (NPs diameter and AC hydrogel mean mesh size) exhibited unexpectedly different roles with regards to the two hindrance contributions involved in mass transport. 3.6. NPs delivery in vivo The hydrogel system described in this paper represents a polymeric formulated network here engineered for multiple NPs delivery, with different and independent kinetic profiles. Biocompatibility was assessed in previous studies by examining different times after gel positioning, in accordance with a translational clinical approach [14]. From in vitro studies (Fig. 3) it is evident that, by changing NPs diameter, it is possible to tune their release rate. In order to further verify the in vitro release kinetics and to prove the reliability of the investigated composite material library, in vivo experiments were carried out too (Fig. 4). The importance of this work resides in the fact that sometimes in vivo conditions are extremely different respect to in vitro ones and hence translation into clinic is not always given for granted [47]. As observable from Fig. 4A, the AC gelling solution was homogenized with NPs solution above the sol–gel transition temperature and then injected into the spinal cord. The gelling solution was injected using a micro-invasive technique [27] and the possibility to direct the polycondensation of the composite material in situ within the target tissue was verified (Fig. 4B). AC1-NPs60 and AC1-NPs130 composite materials were injected in the intrathecal space. Intrathecal drug delivery has been clinically used as an alternative route of drug administration to the systemic and oral ways, expanding the medical options available to treat injured spinal cord. It is proposed as a direct way to release compounds with a single or continuous infusion directly into the CSF at the cord level. Advantages with respect to direct drug injection reside in higher distribution of therapeutic compounds, here able to reach also the inner part of the spinal parenchyma. On the contrary, due to the CSF flow [14], direct injected compounds return quickly in the bloodstream and the BSCB represents again a strong limitation to a diffuse pharmacological treatment of the spinal cord. Here, thanks to the hydrogel acting as carrier, the ability of the delivering system to remain in situ allows releasing active compounds into the CSF and then into the target spinal cord. After 7 days hydrogels were recovered and analyzed with DLS in to order to quantify the percentage of NPs released. The comparison between NPs release studies in vitro and in vivo shows a good qualitative agreement. Indeed, the amount of NPs released after 7 days (Fig. 4C) is strictly dependent on their diameter, in particular is higher for NPs 60 nm, reaching about 60% of the loaded NPs. Hence, in vivo studies demonstrated the reliability of in vitro studies together with the possibility to tune the release rates of NPs depending on NPs diameter. 4. Conclusions It is well known that drug release kinetics have dramatic effects on the efficacy of therapies, and this consideration strongly applies also to SCI treatments, where a multi-target approach with an
extended therapeutic window is required. Consequently, among the strategies for SCI repair, one of the main priorities in drug delivery systems design is to obtain tunable release rates, with fine control over them. In this framework, polymeric NPs are good candidates for controlled release of different drugs, but their potentiality can be relevantly improved loading them within a hydrogel that offers control over their release directly in situ, within the damaged tissue. Here, a hydrogel-based system already tested in SCI repair strategies, was loaded with different sized NPs and the diffusion through hydrogel pores was studied both in vitro and in vivo on a reliable translational animal model. The composite hydrogel–NPs system efficacy can be improved and finely tuned, thus giving origin to a material library capable of satisfying different therapeutic needs with respect to desired release kinetics. In defining such a material library, the influence of both NPs diameter and AC mean mesh size was taken into account and their role was precisely understood, underlying the key role of NPs diameter, being here more influent than hydrogel mean mesh size.
Acknowledgments This study was supported by Fondazione Cariplo, grant no. 2010/0639. The authors would like to thank Dr. Hua Wu (ETH Zurich) and Prof. Maurizio Masi (Politecnico di Milano) for fruitful discussions and Dr. Bertrando Sacchi for having provided ultrapure pharmaceutical grade carbomer.
Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.colsurfb. 2013.02.046.
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Colloids and Surfaces B: Biointerfaces journal homepage: www.elsevier.com/locate/colsurfb
Tunable hydrogel—Nanoparticles release system for sustained combination therapies in the spinal cord Filippo Rossi a,b , Raffaele Ferrari a , Simonetta Papa b , Davide Moscatelli a , Tommaso Casalini a , Gianluigi Forloni b , Giuseppe Perale a,b,∗ , Pietro Veglianese b a b
Politecnico di Milano, Dipartimento di Chimica, Materiali e Ingegneria Chimica “Giulio Natta”, via Mancinelli 7, 20131 Milano, Italy Istituto di Ricerche Farmacologiche Mario Negri-IRCCS, Dipartimento di Neuroscienze, via La Masa 19, 20156 Milano, Italy
a r t i c l e
i n f o
Article history: Received 28 December 2012 Received in revised form 26 February 2013 Accepted 27 February 2013 Available online 14 March 2013 Keywords: Drug delivery Hydrogel Polymeric nanoparticles Release system Spinal cord
a b s t r a c t Poly(methyl methacrylate) (PMMA) nanoparticles (NPs) were prepared by emulsion free radical polymerization. NPs with controlled dimension, as monitored by dynamic light scattering (DLS) and transmission electron microscopy (TEM), were produced by changing experimental parameters, such as the amount of emulsifier and the monomer feeding mode (batch or semi-batch). Then, different sized NPs (60, 80 and 130 nm) were loaded in polysaccharide-polyacrylic acid based hydrogels, cross-linked by covalent ester bonds between polyacrylic acid (PAA) and agarose chains, with different pore sizes (30, 60, 90 nm). The characteristics of the resulting composite hydrogel-NPs system were firstly studied in terms of rheological properties and ability to release Rhodamine B that presents steric hindrance similar to many neuroprotective agents used in spinal cord injury (SCI) repair. Then, diffusion-controlled release of different sized NPs from different entangled hydrogels was investigated in vitro and correlated with NPs diameter and hydrogel mean mesh size, showing different hindrances to the diffusion pathways. Release experiments and diffusion studies, rationalized by mathematical modeling and verified in vivo, allowed to build a material library able to satisfy different medical drug delivery needs. © 2013 Elsevier B.V. All rights reserved.
1. Introduction Traumatic spinal cord injury (SCI) is a dramatic condition that results in profound and life-lasting disability for the patient, with consequent reduction of the quality of life. Its incidence in the EU ranges between 28 and 55 cases per million people, with about 11.000 new cases reported every year [1–3] in the sole Europe. The lesion mainly results from a contusive, compressive or stretch injury of the cord. However, most of the post-traumatic degeneration of the tissue is due to the so-called secondary injury, which is known to be a multifactorial process [4–7]. In the last decades the primary interest of the scientific community has been pointing toward the mechanisms of secondary injury in order to minimize its pathological consequences [8,9]. Indeed, a growing number of pharmacological treatments have been proposed as potential therapeutic approaches, but nowadays none of them showed evident therapeutic efficacy when translated to humans [10]. Probably this is due to the fact that several therapeutic approaches aim to counteract only one single effect, while different therapeutic targets
∗ Corresponding author at: Dipartimento di Chimica, Materiali e Ingegneria Chimica “Giulio Natta”, Politecnico di Milano, via Mancinelli 7, 20131 Milano, Italy. Tel.: +39 02 2399 3145; fax: +39 02 2399 3180. E-mail address: [email protected] (G. Perale). 0927-7765/$ – see front matter © 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.colsurfb.2013.02.046
should be considered at the same time: multifunctional therapies directed to counteract multiple injury mechanisms, trying to combine both neuroprotective and neuroregenerative agents, seem to be extremely promising [3,4,11,12]. Hence, the ability to guarantee different controlled drug delivery kinetics using neuroprotective agents in the first days and neuroregenerative in the following ones could be an efficacious therapeutic strategy [13–15]. In this framework, recent advances in polymer science have provided a huge amount of innovations, underlining the increasing importance of macromolecules in controlled release applications [16–20]. Particularly NPs, thanks to their versatility in terms of size, potential surface and hydrophilic/lipophilic characteristics, lead relevant advantages in drug delivery by increasing the selectivity of drugs and by controlling their release during time [21–25]. Hence, a good strategy could be to load drugs within NPs in order to deliver hydrophobic drugs into injured spinal cord for therapeutic purposes [8,26]. Direct injection of colloidal NPs suspension is one attractive alternative, but several concerns arise: injected NPs very often leave the zone of injection as they are not confined by any support, and easily extravasate into the circulatory torrent, migrating all over the body to the liver and the spleen, or toward an uncertain faith [8,27]. According to these critical issues, nanostructured hydrogels could be reliable NPs carriers, due to their ability to control release rates in situ [13,28,29].
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In this framework, we investigated the potential use of PMMA NPs loaded within hydrogel networks, hence forming composite materials [1], for drug delivery purposes in SCI repair strategies. In particular it should be underlined that NPs directly injected within the spinal cord are quickly (i.e. few hours) uptaken by macrophages [30,31]. Indeed, the controlled release of drug loaded NPs from hydrogels could help to both sustain a selective and effective release of drugs and also delay the cell uptake mechanism. The chosen hydrogel library, specifically developed for SCI repair strategies, was obtained by synthesis from statistical block polycondensation between agarose and carbomer 974P, together with cross-linkers (being a library of hydrogels, they differ by means of components mutual ratios and hereafter they are briefly termed as AC, acronymic of their main components) [27,32]. In previous works we observed that our hydrogels can remain localized at the site of injection [27], showing high biocompatibility and good ability to provide short term delivery for neuroprotective molecules together with long term for neuroregenerative ones [14]. PMMA NPs were synthesized through emulsion free radical polymerization process, which ensures a very narrow particle size distribution and an easy control over final average particle size [33,34]. Then, after AC synthesis but before sol/gel transition takes place, NPs were physically entrapped within the AC networks forming ACNPs composite materials. Since PMMA based NPs are not easy biodegradable, they can be traced both in in vitro and in vivo biological studies over a long period of time without side effects due to the biodegradation process. Obtained composite materials were characterized from a physical-chemical point of view and compared with neat AC hydrogels, already studied in previous works [32]. NPs release from AC hydrogels was studied with the aim to develop a new library of hydrogel-nanoparticles based materials able to provide different release rates, tuning independently both hydrogel mean mesh size (30, 60, 90 nm) and NPs diameter (60, 80, 130 nm): ability to provide different release kinetics was demonstrated and also verified in vivo in a spinal cord mouse model. 2. Experimental 2.1. Materials For NPs synthesis, methyl methacrylate (MMA, 99% purity, Sigma–Aldrich, Germany), polyoxyethylene sorbitan monooleate (Tween80, Sigma–Aldrich, Germany) and potassium persulfate (KPS, >99% purity, Sigma–Aldrich, Germany) were all used as received. For hydrogel synthesis: cross-linked polyacrylic acid carbomer 974P was purchased by Fagron (The Netherlands), phosphate buffered saline solution (PBS), propylene glycol, glycerol and sodium hydroxide were purchased by Sigma–Aldrich (Germany) while agarose from Invitrogen (Carlsbad, CA, USA) and all were used as received. Rhodamine B base (RhB) (Sb sensitivity <0.1 g mL−1 , Carlo Erba reagents, Italy) was used for release experiments. 2.2. NPs synthesis Emulsion free radical polymerization reaction of MMA was carried out in batch and semi-batch conditions, as already described [35], using a 100 mL three necked glass flask equipped with a reflux condenser; temperature were controlled with an external oil bath. Different amounts of Tween80 were added to 100 mL of deionized water and the solution was heated up to 80 ◦ C. Vacuum-nitrogen cycles have been used to make and to maintain inert the reaction atmosphere. The initiator, KPS, in concentration of 1% wtKPS /wtMMA was added to the purged solution. Next, 5 g of MMA was injected with a flow rate of 3 mL/h using a syringe pump (Model NE-300, New Era Pump System, Farmingdale, NY, US) for semi-batch
reaction, while for batch reaction 5 g of MMA were all injected before the initiator. In both cases reaction was run for 3 h. Final monomer conversion was measured through gravimetric measurement and calculated as higher than 98.5%. Different particle sizes were obtained varying the emulsifier-monomer ratio and the feeding condition (batch and semi-batch). In particular, NPs of 60 and 80 nm were produced using semibatch conditions with 30% and 20% wtTween80 /wtMMA ratio, while NPs of 130 nm were obtained with a 15% wtTween80 /wtMMA ratio in batch conditions. RhB loaded NPs of 130 nm were produced using the same experimental condition; where MMA and RhB (5 mg) solutions were fed into the reactor. RhB concentration was measured by spectrofluorimetry at excitation (ex) and emission (em) wavelengths of 552 and 585 nm, respectively, using a Tecan Safire plate reader (Durham, NC). Percent encapsulation efficiency (% EE) was calculated based on the following equation: % EE =
Drug entrapped in NPs · 100 Initial amount of drug added
(1)
2.3. Characterization of NPs Final particle size has been determined by DLS (Malvern, Zetanano ZS, US); analyses were performed in triplicate and reported data are the average of three runs where the standard deviation was always below 5%. Dimension and morphology of produced NPs were also confirmed by transmission electron microscopy (TEM; using a EFTEM Leo 912AB, at 80 kV, by Karl Zeiss, Jena, Germany). Samples were prepared placing 5 L drop of NPs dispersion on a Formvar/carbon-coated copper grid and dried overnight. Digital images were acquired by a charge coupled device (CCD; Esi Vision Proscan camera). 2.4. Hydrogel synthesis Polymeric solution was achieved by mixing polymer powders of agarose and carbomer 974P in PBS (Phosphate Buffer Saline solution) in batch condition, adding a mixture of cross-linking agents made of propylene glycol and glycerol along with NaOH 1 N (reaction pH was kept neutral) [14,32]. The gelation onset was achieved by means of electromagnetic stimulation (500 W irradiated power) heating in ratio of 1 min per 10 ml of polymeric solution at 80 ◦ C. Microwave-enhanced chemistry is based on the efficient heating of materials by “microwave dielectric heating” effects [36]. This phenomenon is dependent on the ability of a specific material (solvent or reagent) to absorb microwave energy and convert it into heat. This efficient method could allow gel forming in less than 10 min, avoiding prolonged heating that could damage the structures of the components and the compounds loaded. Here, carbomer carboxyl groups constitute the main cross-linking sites to be reacted with hydroxyl groups from agarose, altogether giving rise to the three dimensional matrix. Propylene glycol and glycerol take part in the condensation reaction and form local interconnections through their own hydroxyl groups [14]. They also play a physical role, increasing system viscosity and thus enhancing tight macromers cross-linking. AC formulations are numbered according with previous works [35] and their detailed compositions are presented in Supplementary Material. The differences in these three formulations are due to the different presence of cross-linkers, resulting in different mesh size values: nominal 90 nm for AC1, 60 nm for AC3 and 30 nm for AC6 [37]. By varying the amounts and the functionality, f in Flory notation, of the cross-linkers we can synthesize more or less entangled polymeric networks with high impact on physico-chemical properties [37]. Thus, acting on microchemistry of hydrogel library and following a multi scale approach, it was possible to manage macro properties.
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2.5. Rheology Rheological analyses were performed, on AC-NPs systems, at 37 ◦ C, using a Rheometric Scientific ARES (TA Instruments, New Castle, DE, USA) equipped with parallel plates of 30 mm of diameter and a 4 mm gap between. Oscillatory responses, dynamic frequency sweep tests (Gˇı, elastic modulus, and Gˇıˇı, loss/viscous modulus) were determined at low strain values over the frequency range 0.1–500 rad/s [37]. Thixotropic behavior of the samples was also investigated and shear strain and viscosity, as a function of shear rate, were evaluated too [32]. To evaluate the role of hydrogel we performed rheological studies with AC1, AC3 and AC6 loaded with NPs (80 nm), here briefly named AC1-NPs80, AC3-NPs80 and AC6NPs80. The role of NPs diameter was studied comparing AC1-NPs80 with AC1-NPs130. 2.6. Rhodamine B release from AC hydrogel and AC-NPs composite system RhB was chosen mainly because of its steric hindrance, similar to many small drugs, and for its absorbance that makes it easily detectable by UV spectroscopy [11]. It was loaded within (i) AC1 hydrogel network and (ii) within NPs (130 nm) then loaded in AC1 hydrogels (AC1-NPs130) to compare the release ability of these two systems. (i) RhB aqueous solution (0.1 mg/mL) was mixed with AC1 gelling solution, above sol/gel transition to allow good solute dispersion within polymeric network; gelation took place in steel cylinders (0.5 cm3 , d = 1.1 cm). (ii) NPs (loaded with RhB as explained in Section 2.2) solution was mixed with AC1 gelling solution, above sol/gel transition to allow good solute dispersion within polymeric network. In order to allow good comparison between these two systems, RhB concentration was kept constant in both of them. Gelation took place in steel cylinders (0.5 cm3 , d = 1.1 cm). Then, three samples for each condition were put in excess of cerebrospinal fluid (CSF) and aliquots were collected at defined time points, while the sample volume replaced by fresh CSF, in order to avoid mass-transfer equilibrium between the gel and the surrounding solution. Percentage of RhB released was measured by UV spectroscopy [38]. 2.7. NPs Release from AC hydrogels and mathematical modeling AC1, AC3 and AC6 hydrogels were loaded with different PMMA NPs, in terms of diameter (60, 80, 130 nm), to assess their in vitro diffusion-controlled delivery. Composite materials were briefly named as following: AC1-NPs60 for AC1 hydrogel loaded with 60 nm diameter NPs etc.. . . During AC hydrogels cooling phase, slightly above 37 ◦ C, AC samples were mixed with the chosen NPs solutions at 1% monomer concentration: this procedure allows solute loading still during sol state, i.e. before sol/gel transition [14]. The loaded AC-NPs systems were cast in standard plastic 48well cell culture plates, 0.5 mL/well (diameter = 1.1 cm). Complete swelling was achieved leaving samples in a NPs isotonic solution within a standard incubator at 37 ◦ C and 5% CO2 atmosphere, overnight. This procedure is necessary to allow swelling without having any NPs release. Swelling equilibrium was quickly reached in the first hour and, on the other hand, stability of this material was quite high: it remained stable for weeks before significant degradation, by hydrolysis, took place [37]. Hence, as characteristic time is much larger than experimental length, degradation reactions can be taken as fully neglectable in this framework. Therefore, NPs flux is controlled only by concentration gradient and diffusion can be considered as Fickian. Hydrogels were then placed in CSF filled wells (5 mL volume each). Aliquots were collected at defined time points, while the sample volume replaced by fresh CSF, in order to avoid mass-transfer equilibrium between the gel and the
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surrounding solution. Samples collected were analyzed to assess the NPs release percentage using DLS and full details are contained in Supplementary Material. In addition, to fully recapitulate this AC-NPs composite material library, NPs diffusivity (D) was evaluated with a mathematical model based on mass balances, i.e. on fundamental conservation laws [32,39]. Diffusion is described through the second Fick law with a 1-dimensional model in a cylindrical geometry, as shown in Eq. (2). This particular hydrogel shape was chosen to best fit the average cystic cavity present inside an injured spinal cord [1,39]. Here radius (r) is the characteristic dimension for the investigated transport phenomenon. Therefore, the above-mentioned increase takes place due to the material flux, which takes place at the CSF/hydrogel surface. Eqs (6) and (7) represent the boundary conditions for the left and the right border, respectively. The first one implies profile symmetry at the center (that is, with respect to cylinder axis), while the second one represents the equivalence between the material diffusive fluxes at the CSF/hydrogel surface. ∂CG 1 ∂ =D· · · r ∂r ∂t VS ·
r·
∂CG ∂r
∂CS = kC · Sexc · (CG − CS ) ∂t
CS (t = 0) = 0
−D ·
mG,0 VG
=0 r=0
(3) (4)
CG (t = 0) = CG,0 = ∂CG ∂r
(2)
(5)
(6)
∂CG ∂r
= kC · (CG − CS )
(7)
r=R
The two mass balance equations involve the mean NPs concentration within the hydrogel (CG ), the mean NPs concentration in the outer solution (CS ), the volume of the solution (VS ), the hydrogel volume (VG ), the NPs mass present inside the matrix (mG ) and the exchange interfacial surface (Sexc ), i.e. the boundary surface between gel and surrounding solution (which, simplifying, can be here considered as being only the side surface). Finally, D represents the diffusion coefficient and kC , the mass transfer coefficient. The mass transfer coefficient is computed through Sherwood number obtained by means of penetration theory [40]: Sh =
8 kC · 2r = D
(8)
2.8. Animals and their care C57BL/6 mice were used for this study. Procedures involving animals and their care were conducted in accordance with the institutional guidelines, which are in compliance with Italian national (D.L. no. 116, G.U. suppl. 40, Feb. 18, 1992, Circolare No.8, G.U., 14 luglio 1994) along with international laws and policies (EEC Council Directive 86/609, OJL 358, 1 DEC.12, 1987; NIH guide for the Care and use of Laboratory Animals, U.S. National Research Council, 1996). All animals were housed under standard conditions (22 ± 1 ◦ C, 60% relative humidity, 12 h light/dark schedule), one per cage, with free access to food (Altromin MT, Rieper, Vandoies (BZ), Italy) and water. 2.9. Surgery To obtain a reliable translational study and to investigate different NPs release possibilities in vivo, twelve mice were enroled for the surgery and randomly divided into 2 groups of 6 animals each:
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Table 1 Characteristics of PMMA based NPs. Nominal diameter [nm]
Diameter [nm]
PSD
-potential [mV]
60 80 130
63.80 81.52 131.9
0.063 0.062 0.070
−27.1 −22.8 −23.3
animals received AC1-NPs60 and AC1-NPs130 respectively. The NPs concentration loaded is the same used for in vitro studies. All surgery procedures were performed under deep anesthesia by 2.5% Avertin (2,2,2-tribromoethanol in 2-methyl-2-butanol, 400 mg/kg, i.p., Sigma–Aldrich, Germany). Thirty minutes before undergoing surgery, animals received an injection of ampicillin (0.03 mg/kg, subcutaneous, Roche, Italy) and buprenorphine (0.05 mg/kg, subcutaneous, Sigma–Aldrich, Germany). The back of the animal was shaved at dorsal level and a cutaneous incision (3 cm) was performed to expose the backbone. T11 and T12 vertebrae were identified and exposed by separation of dorsal and intervertebral muscles. Animals were then placed in a Cunningham Spinal Cord Adaptor (Stoelting, Ireland) mounted on a stereotaxic frame, and laminectomy of the T11 vertebra was done to uncover thoracic spinal cord. Subsequently, a gelling solution (6 L) was placed with a micropipette on top of the cord in the subdural space [14]. After gel positioning, dorsal muscles were juxtaposed using absorbable sutures and the skin sutured and disinfected. After surgery animals were kept on a warm pad for 30 min and then placed in separated cages for recovery. After seven days animals were anesthetized with Equithesin (1% Phenobarbital, 4% chloral hydrate, 30 mL/10 g, i.p.) and transcardially perfused with 20 mL saline and subsequently with 50 mL of sodium phosphate-buffered 4% formaldehyde solution. Spinal cords were rapidly removed and the hydrogels recovered and analyzed with DLS to evaluate the percentage of NPs present within them. 2.10. Statistical analysis Where applicable, experimental data were analyzed using Analysis of Variance (ANOVA). Statistical significance was set to p value <0.05. Results are presented as mean value ± standard deviation. 3. Results and discussion 3.1. Nanoparticles synthesis As far as emulsion free radical polymerization is a wellestablished process, it is possible to tune the final NPs characteristics by changing process parameters such as the amount of emulsifier, the reaction temperature and the feeding mode [41,42]. It is worth noticing that emulsion polymerization ensures the final biocompatibility of the NPs suspensions: indeed, the reaction is carried out in water without using any solvent and adopting a FDA approved emulsifier such as Tween80 [34]. Beside biocompatibility, this emulsifier was chosen, despite not being as efficient as ionic surfactants (like e.g. sodium dodecyl sulfate), because it does not introduce any charge effects that might affect the NPs release from the hydrogel matrix. Finally, this technique allows the synthesis of polymeric NPs with narrow particle size distribution (PSD), necessary to discriminate NPs with relatively similar particle size. Obtained results are shown in Table 1, where NPs were produced with the process parameter reported in the Experimental Section. As expected, NPs obtained with different amount of emulsifier and feeding mode are different in size. In particular, batch operation mode leads to larger NPs, namely 130 nm, while smaller NPs are produced adopting semi-batch process. Accordingly,
maintaining semi-batch condition, a lower amount of emulsifier leads to bigger NPs. Moreover, PSD is very narrow, with values lower than 0.1 (Table 1), as observable from the DLS and TEM images of the NPs, presented in Supplementary Material, which shows how NPs present similar sizes; some aggregates are detected but they are due to the sample preparation technique which leads to aggregates formation, anyway each single particle is very well visible. As a final remark, -potential data are in the expected range of NPs produced using steric emulsifier as Tween80, where the charge of the NPs come only from the initiator, KPS [43].
3.2. Rheology The storage modulus, G , was compared between unloaded gel samples (AC1) and NPs loaded ones (AC1-NPs80) to investigate the effect of NPs on the rheological behavior of these composites (Fig. 1A). In previous works gel storage modulus (G ) was found to be approximately one order of magnitude higher than the loss modulus (G ), indicating an elastic rather than viscous material [37]. Furthermore, both G and G of each gel were essentially independent from frequency over the entire investigated range, thus indicating dominant viscoelastic relaxations of networks at lower frequencies. Such rheological behavior matches the characteristic signature of a solid-like gel, confirming this nature for AC type gels (see Supplementary Material). These values are in complete accordance with literature and, in particular, with the requirements for SCI repair applications [1]. The presence of NPs does not affect the rheological properties of the composite system and also in this case G trends are essentially independent from frequency over the range investigated. In comparing neat systems with those containing NPs, G increase is evident: specifically, the addition of NPs immediately results in a 3-fold increase in G over neat AC gels, thus suggesting a stiffer and more elastic material. The increased elastic modulus could be attributed to interactions between AC network and NPs. In particular electrostatic interactions between negatively charged AC network and negatively charged PMMA NPs increase the chain mobility, hence increasing the whole system elastic properties. The consistent and monotonic enhancement in hydrogel mechanical properties also suggests that NPs are well dispersed within the polymeric matrix. In addition, AC hydrogels do not show any loss in elasticity upon addition of NPs, suggesting that the underlying polymer network remains intact upon NPs addition and that it does not change the morphology and the shape of loaded NPs. This assumption is further corroborated by evidences from DLS (see Supplementary Material, Table S.2). Following the oscillation frequency sweep, the thixotropic loop method with a flow shear-rate sweep was performed on both neat and NPs-loaded AC gels results are presented in Fig. 1B: the typical area visible from thixotropic loop method corresponds to the energy needed to breakdown the structural stability of the material and the time necessary for the rearrangement of the material after being sheared. Once the structural effect on rheological properties was confirmed, the NPs loading contribution was considered: hysteresis loop areas are bigger for both NPs-loaded hydrogels with respect to neat ones. The presence of NPs, indeed, makes the gels more stable and leads to higher mechanical properties, in agreement with dynamic frequency sweep tests. In summary, NPs loaded composite materials, compared with neat ones, show faster sol/gel transition and this effect could be a relevant improvement for injectability [44]. Similar results were obtained also with AC3 and AC6 gels, here not presented for bare sake images clarity (see Supplementary Material, Fig. S.3). The same investigations were also performed considering NPs of all different sizes: as predictable, gel elastic modulus G shows a high dependence on NPs diameter, increasing with NPs size increase (see Supplementary Material).
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Fig. 1. Rheological analyses: (A) dynamic frequency sweep test of composite material at 37 ◦ C with small oscillatory shear in the linear viscoelastic regime: (A) G ( AC1, 䊉 AC1-NPs). (B) Thixotropic loop for: AC1 () and AC1-NPs (䊉).
3.3. Rhodamine B release Once proved that the AC-NPs composite systems maintain adequate mechanical/structural properties, RhB release was investigated from both neat AC and AC-NPs hydrogels. Firstly, the in vitro drug release kinetics of RhB, loaded in neat (a) AC1 hydrogel and in (b) NPs (130 nm) then loaded in AC1 hydrogels (AC1-NPs130), were deeply investigated and compared. The importance of this study resides in the necessity to prove NPs ability to release drugs with proper kinetic on one side and, on the other, to verify the advantages behind the combination of hydrogels together with NPs in a composite release system. As said, RhB was chosen for this investigation because its steric hindrance resembles many corticosteroids and anti-inflammatory drugs (for example methylprednisolone and estradiol), commonly used in pharmacotherapy to hinder SCI secondary effects and CNS damages [8]. RhB release profiles from NPs and AC1-NPs130 systems are compared in Fig. 2: from Fig. 2A it is well visible that RhB release profile from AC1 samples is very fast and it is almost completed in the first 12 h, while RhB release from AC1-NPs130 system is prolonged for a much longer period; from Fig. 2B it is also evident that loading NPs within hydrogels is hence the most promising solution. There the influence of the system in delivering RhB was investigated plotting release percentage against time square root. A linear plot is indicative of Fickian diffusion and the y-axis intercept value an indication of
burst release, where it is well known that an ideal controlled release system should present linear trend during time and its y-axis intercept equal to zero [45]. RhB directly loaded within AC hydrogels shows a linear trend only in the first hours and then a plateau trend is visible. Moreover, in this case the burst release contribution is high (about 40%), underlining the expected poor ability to control release of small drugs [14]. On the other hand, loading NPs, that contain RhB, within AC hydrogels increases performances in terms of pure diffusive mass transport and of burst release contribution that results to be extremely small (around 1%). This is due to the fact that AC and NPs represent multiple diffusion barriers to RhB release and so the AC1-NPs130 system can control and sustain better the release of small drugs. Thus, after having verified the increased release ability of small drug from AC1-NPs system, we investigated the possibility to build an AC-NPs material library that can tailor the diffusion during time of NPs and, consequently, of the loaded drugs. 3.4. NPs release from hydrogel We discussed above how a small drug mimetic (RhB) can be released from AC-NPs composite systems and hence the advantages to load NPs within AC hydrogels. We further investigated the possibility to build a composite material library loading NPs with different diameter within hydrogels with different pore size,
Fig. 2. (A) In vitro release profiles of RhB from AC1 hydrogel () and AC1-NPs130 composite system (䊉) in CSF at 37 ◦ C. (B) The slope of RhB releases against the square root time is representing how the delivery system influences Fickian diffusion.
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Fig. 3. (A) Cumulative release of NPs with different diameters (red = 130 nm; blue = 80 nm; black = 60 nm) from AC1 hydrogel. The same trends were also found with AC3 and AC6 formulations here not presented for image clarity; (B) release of NPs (diameter = 80 nm) from AC1 () and AC6 () hydrogels. (C) The slope of NPs (red = 130 nm; blue = 80 nm; black = 60 nm) release from AC1 against the square root of time is representing how NPs diameter influences Fickian diffusion. (D) The slope of NPs release from AC1 (), AC3 () and AC6 () against the square root of time is representing the role of hydrogel mesh size in mass transport. Data points are presented mean ± standard deviation of three different gel cylinders. (E) Effect of NPs diameter and AC hydrogel mean mesh size on NPs diffusion coefficient: D trend exhibits high dependence on NPs diameter (red = 130 nm; blue = 80 nm; black = 60 nm) while AC mean mesh size is less able to influence mass transport through the network pores. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of the article.)
evaluating their release during time. Our previous works [14,32] have demonstrated that solutes physically entrapped within the network leave completely hydrogel matrix with kinetics depending on their hydrodynamic radius: mass flux is, indeed, driven only by concentration gradient and diffusion can be considered as Fickian, with no loss of generality as long as gel degradation characteristic time is much higher than diffusion one. Here, to investigate the influence of particle size on the release kinetics, different sized NPs were loaded within different hydrogel networks at sol state, just before sol/gel transition. In Fig. 3A the release kinetics from AC1 gel formulation with three different NPs diameters are
presented. The same trends were also found with AC3 and AC6 formulations here not presented for bare sake of image clarity (see Supplementary Material, Fig. S.5). The percentage of NPs released from AC-NPs in the first 2 h, less than 30%, is attributed to the high initial concentration gradient and in particular at NPs that were at or near the solvent–hydrogel interface and thus could rapidly diffuse into the supernatant solution. Thereon, NPs were slowly released showing a diameter dependent kinetic: the bigger NPs need more time to be released from the hydrogels than the smaller ones. In Fig. 3B, on the other hand, the release kinetics of the same NPs (80 nm) from different AC formulation (AC1 and AC6) are
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Fig. 4. (A) The AC gelling solution was homogenized with NPs solution (red 130 nm, yellow 60 nm) above the sol–gel transition temperature and then injected into the spinal cord. After 7 days AC-NPs was recovered and analyzed with DLS to quantify the in vivo percentage of NPs with different diameters (60 nm and 130 nm) released from AC1 hydrogel; (B) picture of in situ gelified AC-NPs composite system after 7 days before hydrogel recovering (white rectangle; scale bar = 0.5 mm); (C); percentages of NPs60 and NPs130 released from AC1 hydrogels after 7 days. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of the article.)
reported. AC1 hydrogel generally showed the fastest release when compared to all other formulations. From Fig. 3A and B it is evident that hydrogel mesh sizes is not the limiting factor for drug release while NPs diameter is the real limiting parameter. This is due to the fact that the calculated hydrogel mesh size is an average value [28]: they are synthesized from statistical random walk reaction and so they are not polymeric ordered matrices. Moreover, to better investigate and understand the influence of NPs diameter and AC mesh sizes, release percentage is plotted against time square root (see Fig. 3C and D). NPs 60 and 80 nm showed linear behavior only in the first hours (Fig. 3C), while for NPs 130 nm this is evident for the whole considered timeframe. This difference is due to the initial contribution of burst release that is small for low NPs values and is not present considering NPs 130 nm. Hence, increasing NPs diameter, initial burst release could be limited and a linear trend for all experimental times could be provided. On the opposite, as explained above, AC mesh size role in NPs delivery tuning is less important: from Fig. 4D is visible that, changing mean mesh size, the mass transport is only partially influenced. So, the investigation of parameters roles allows high drug release kinetic tunability according to different potential therapeutic approaches in SCI repair. In particular it should be underlined that the most probable reason behind disappointing results in SCI treatment is due to the fact that SCI is characterized by a temporal development of biochemical pathways of degeneration. Recent research has focused its attention on multifactorial therapies
directed to counteract multiple injury mechanisms, trying to combine both neuroprotective and neuroregenerative agents [12,46]. In particular, as explained above, controlling the release of NPs-drug loaded consents a selective and sustained treatment. 3.5. NPs diffusivity evaluation Mass release data obtained experimentally (Fig. 3A–D) were used to estimate NPs diffusion coefficients. As explained above, the release mechanism could be considered as a pure Fickian diffusion, being concentration driven through AC hydrogel pores and allowing to neglect any contribution of both swelling and degradation phenomena. Fig. 3E shows the dependence of NPs diffusivity upon their diameter and AC mean mesh size. The diffusion coefficient, magnitude order of 10−6 cm2 /s, does not show any high differences among the three AC formulations studied, even if materials are structurally and mechanically different. In particular, it should be noted that NPs with average diameter of 130 and 80 nm have bigger hindrances with respect to the mean AC mesh sizes and so their diffusivity trends are similar: no evident differences can be observed if we consider an iso-NPs diameter line in the graph. On the other hand, if we consider 60 nm NPs, a slope increase is well visible between mesh size values of 30 and 60 nm: this is reasonable because decreasing the entanglement of the matrix (from AC6 to AC3) the diffusion is favored. In general it seems that, as described above, NPs mobility through the pores is affected only
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partially by the surrounding hydrogel inner network conformation and, hence, mean mesh size represents only a partial hindrance to NPs diffusion pathway. In contrast, NPs diffusivity showed high dependence on particle size with a decreasing trend from small diameters to big ones. In summary, NPs diffusivity values are in accordance with expectations deducible from previous works [32], exhibiting lower values respect to the drug directly loaded in neat AC hydrogels, due to their higher hydrodynamic radius. This result is, indeed, a direct consequence of the Stokes-Einstein equation, which describes diffusivity in inverse proportion to hydrodynamic radius of released molecule. Nevertheless, it is evident that the two parameters involved (NPs diameter and AC hydrogel mean mesh size) exhibited unexpectedly different roles with regards to the two hindrance contributions involved in mass transport. 3.6. NPs delivery in vivo The hydrogel system described in this paper represents a polymeric formulated network here engineered for multiple NPs delivery, with different and independent kinetic profiles. Biocompatibility was assessed in previous studies by examining different times after gel positioning, in accordance with a translational clinical approach [14]. From in vitro studies (Fig. 3) it is evident that, by changing NPs diameter, it is possible to tune their release rate. In order to further verify the in vitro release kinetics and to prove the reliability of the investigated composite material library, in vivo experiments were carried out too (Fig. 4). The importance of this work resides in the fact that sometimes in vivo conditions are extremely different respect to in vitro ones and hence translation into clinic is not always given for granted [47]. As observable from Fig. 4A, the AC gelling solution was homogenized with NPs solution above the sol–gel transition temperature and then injected into the spinal cord. The gelling solution was injected using a micro-invasive technique [27] and the possibility to direct the polycondensation of the composite material in situ within the target tissue was verified (Fig. 4B). AC1-NPs60 and AC1-NPs130 composite materials were injected in the intrathecal space. Intrathecal drug delivery has been clinically used as an alternative route of drug administration to the systemic and oral ways, expanding the medical options available to treat injured spinal cord. It is proposed as a direct way to release compounds with a single or continuous infusion directly into the CSF at the cord level. Advantages with respect to direct drug injection reside in higher distribution of therapeutic compounds, here able to reach also the inner part of the spinal parenchyma. On the contrary, due to the CSF flow [14], direct injected compounds return quickly in the bloodstream and the BSCB represents again a strong limitation to a diffuse pharmacological treatment of the spinal cord. Here, thanks to the hydrogel acting as carrier, the ability of the delivering system to remain in situ allows releasing active compounds into the CSF and then into the target spinal cord. After 7 days hydrogels were recovered and analyzed with DLS in to order to quantify the percentage of NPs released. The comparison between NPs release studies in vitro and in vivo shows a good qualitative agreement. Indeed, the amount of NPs released after 7 days (Fig. 4C) is strictly dependent on their diameter, in particular is higher for NPs 60 nm, reaching about 60% of the loaded NPs. Hence, in vivo studies demonstrated the reliability of in vitro studies together with the possibility to tune the release rates of NPs depending on NPs diameter. 4. Conclusions It is well known that drug release kinetics have dramatic effects on the efficacy of therapies, and this consideration strongly applies also to SCI treatments, where a multi-target approach with an
extended therapeutic window is required. Consequently, among the strategies for SCI repair, one of the main priorities in drug delivery systems design is to obtain tunable release rates, with fine control over them. In this framework, polymeric NPs are good candidates for controlled release of different drugs, but their potentiality can be relevantly improved loading them within a hydrogel that offers control over their release directly in situ, within the damaged tissue. Here, a hydrogel-based system already tested in SCI repair strategies, was loaded with different sized NPs and the diffusion through hydrogel pores was studied both in vitro and in vivo on a reliable translational animal model. The composite hydrogel–NPs system efficacy can be improved and finely tuned, thus giving origin to a material library capable of satisfying different therapeutic needs with respect to desired release kinetics. In defining such a material library, the influence of both NPs diameter and AC mean mesh size was taken into account and their role was precisely understood, underlying the key role of NPs diameter, being here more influent than hydrogel mean mesh size.
Acknowledgments This study was supported by Fondazione Cariplo, grant no. 2010/0639. The authors would like to thank Dr. Hua Wu (ETH Zurich) and Prof. Maurizio Masi (Politecnico di Milano) for fruitful discussions and Dr. Bertrando Sacchi for having provided ultrapure pharmaceutical grade carbomer.
Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.colsurfb. 2013.02.046.
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