J Mater Sci: Mater Med (2014) 25:383–390 DOI 10.1007/s10856-013-5080-5

Design of electrospayed non-spherical poly (L-lactide-co-glicolide) microdevices for sustained drug delivery Laura Mayol • Assunta Borzacchiello • Vincenzo Guarino • Carla Serri • Marco Biondi Luigi Ambrosio

•

Received: 29 July 2013 / Accepted: 21 October 2013 / Published online: 12 November 2013 Ó Springer Science+Business Media New York 2013

Abstract Polymer chain entanglements in organic solvents can be considered a key parameter in the formation of non-spherical beads when electrospraying is employed. The shape of micro/nanometric drug delivery systems plays a major role since it can affect circulation, extravasation, distribution and in vivo clearance of the devices. In this frame, we investigated the influence of polymer processing parameters on the design of polylactic-co-glycolic acid nonspherical microdevices loaded with triamcinolone acetonide (TrA), a sparingly water soluble corticosteroid, prepared by electrospraying technique through a one-step process. In particular, we verified that the formation of non-spherical MDs is related to the presence of entanglements among polymer chains to select the optimal solution to be sprayed. The addition of TrA did not substantially affect the particle morphology in terms of size, size distribution and circularity at all the tested drug loadings. Furthermore, the drug could be released for a prolonged period, with controlled and reproducible kinetics for over 3 weeks. The mathematical modeling of release profiles highlighted that the release is mainly

Laura Mayol, Assunta Borzacchiello, and Vincenzo Guarino have contributed equally to this study. L. Mayol  C. Serri  M. Biondi (&) Dipartimento di Farmacia, Universita` di Napoli Federico II, Via D. Montesano 49, Naples, Italy e-mail: [email protected] L. Mayol  A. Borzacchiello  M. Biondi Interdisciplinary Research Centre on Biomaterials (CRIB), Universita` di Napoli Federico II, P.le Tecchio, 80, Naples, Italy A. Borzacchiello  V. Guarino  L. Ambrosio Istituto per i Materiali Compositi e Biomedici IMCB and DSCTM-CNR, P.le Tecchio, 80, Naples, Italy

driven by degradation, at a higher extent in the case of low drug loading.

1 Introduction Among the technological features of drug delivery systems (DDS), geometry plays a major role. Indeed, it has been seen that non-spherical shape affects circulation, extravasation and distribution of nanosystems in vivo. For example, disc-shaped nanoparticles have demonstrated higher in vivo specificity towards endothelial cells compared to spherical particles of similar size [1]. Moreover, elongated nanoparticles more easily align with blood flow, thus favoring their transport in tissues and organs [2, 3], and also particle clearance and internalisation are enhanced/ accelerated in the case of devices with high aspect ratio [4]. Finally, flow properties of microsized powders intended for inhalation strongly depend on particle aerodynamic diameter, which is affected not only by the geometric size but also by porosity and shape. Different flow properties of microsized respirable powders can, in turn, influence their final therapeutic effect (local or systemic) [5]. Indeed, properly engineered powders for inhalation must be in the 1–5 lm aerodynamic diameter range for deep lung penetration and to exert a systemic effect [6], while bigger particles must be trapped within superior airways for a localised action [7]. Furthermore, shape is a major property of biodegradable microparticles since it affects their degradation rate and hence drug release kinetics [8–10]. Recent literature shows that polymer chain entanglements play a key role in the fabrication of nanoparticles by the electrospraying (ES) process. ES has gained widespread attention over the last decade because of its simplicity of use, versatility, and the ability to engineer nanoparticles with

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controlled size and shape at the micro- and nanoscale. The ease of operation and cost-effectiveness of this technique, along with the possibility to spray a wide range of polymers, have led to a strong interest also in the production of drug delivery devices [11]. For example, electrosprayed micro/ nanoparticles for the controlled release of proteins, nucleic acids, antitumor drugs and antibiotics have been fabricated [12–15]. Moreover, their combination with electrospun nanofibres by simultaneous or sequential deposition opens the way towards the development of new interesting applications in tumor targeting and oral delivery [16]. Hence, ES technique is particularly attractive for the pharmaceutical industry due to its ease of scalability compared to the currently used fabrication methods of micro- and nanoparticles, such as emulsion, nanoprecipitation and coacervation [17, 18]. Recent literature shows that the versatility of ES technique allows to obtain non-spherical micro/nanoparticles with a controlled shape by properly tailoring the chemico-physical properties of the solution(s) to be sprayed. In particular, it has been observed that polymeric chain entanglements play a key role in ES process. Indeed, by properly choosing polymer molecular weight and concentration, it is possible to discriminate between the formation of fibers or particles and, in the latter case, chain entanglements can strongly affect size and morphology of the obtained particles, thus offering tunable release kinetics suitable for different DDS [19, 20]. In this frame, we aimed at investigating the role of polymer chemical/physical parameters on the formation of electrosprayed PLGA nSMDs for the sustained release of triamcinolone acetonide (TrA), chosen as a model drug, and the possibility to load the drug in a one step process. Polylactic-coglycolic acid (PLGA) was chosen because of its well-known biocompatibility and hydrolysis-driven biodegradation. PLGA has been widely employed in the field of controlled drug delivery to load a wide array of drugs, from highly hydrophilic to very lipophilic [21]. TrA, a sparingly water-soluble corticosteroid, was chosen as a model drug. Rheological studies were carried out on solutions of PLGA in ethyl acetate (EA). The morphology of the obtained non-spherical microdevices (nSMDs) examined by scanning electron microscopy (SEM), and TrA release kinetics were studied by spectrophotometric analysis. A mathematical model, developed by Corrigan et al. [22], taking into account drug desorption and polymer degradation, was also used to investigate on the key phenomena controlling TrA release from nSMDs.

2 Materials and methods 2.1 Materials Equimolar uncapped poly(D,L-lactide-co-glycolide) (PLGA) (Resomer RG502H, inherent viscosity in chloroform at

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25 °C: 0.16–0.24 dl/g) was purchased from Boehringer Ingelheim (Ingelheim, Germany). EA and the salts for phosphate buffered solution (PBS) were obtained from SigmaAldrich (USA). Distilled water from Milli-Q (Millipore, USA) was used. TrA was kindly supplied by Fisiopharma (Palomonte, Italy).

2.2 Rheological properties of PLGA solution PLGA solutions in EA were prepared at different concentrations in (0.01–5 % w/v) for a rheological analysis aimed to select the appropriate concentration to be sprayed. The rheological properties of PLGA solutions were analysed through a stress controlled rotational rheometer (Gemini Bohlin Instruments, UK), using coaxial cylinders geometry and parallel plate geometry (PP30 cell), at controlled temperature of 25 °C. Steady-state measurements were performed to evaluate solution equilibrium and the dependence of viscosity upon shear rate, i.e. the so called flow curves in the 0.1–100 s-1 range. Small-amplitude oscillatory shear experiments were performed to evaluate the time-dependent response and the linear viscoelastic properties (G0 and G00 ). The frequency was in the 0.1–10 Hz range. In a dynamic test, the material is subjected to a sinusoidal shear strain: c ¼ c0 sinðxtÞ;

ð1Þ

where c0 is the shear strain amplitude, c the oscillation frequency (which can be also expressed as 2pf, where f is the frequency in Hz) and t the time. The mechanical response, expressed as shear stress s, is intermediate between an ideal pure elastic solid (obeying Hooke’s law) and an ideal pure viscous fluid (obeying Newton’s law) and, therefore, it is out of phase with respect to the imposed deformation as expressed by: s ¼ G0 ðxÞc0 sinðxtÞ þ G00 ðxÞc0 cosðxtÞ

ð2Þ

where G0 (x) is the storage or elastic modulus and G00 (x)the loss or viscous modulus. G0 (x) gives information about the elasticity, i.e. the energy stored in the material during deformation, whereas G00 (x) describes the viscous component, i.e. the energy dissipated as heat. Since viscoelastic properties strongly depend upon the time-scale of observation, the dependence of G0 (x) and G00 (x) upon the frequency, the so called mechanical spectrum will be reported. In order to identify the linear viscoelastic response range of the tested solutions, preliminary strain sweep tests were performed on the samples, at the oscillation frequency of 1 Hz. The tests were repeated at least three times on each sample.

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2.3 Fabrication of nSMDs Electrospraying was performed with an automated equipment (Nanon01, MECC, Japan) by transferring polymeric solutions into a 5-mL plastic syringe (BD syringe) and continuously forcing it through the spraying nozzle using a programmed pump. Flow rate values of 0.2 and 0.4 mL/h were chosen and two different voltage values (15 and 18 kV) were applied to the spraying nozzle to initiate atomisation of the charged droplets. Particles were collected onto an aluminum foil (5 9 5 cm). The distance between electrodes was 10 or 13 cm to ensure the right balance between electric forces and evaporation mechanisms during the particle deposition. Temperature and relative water content were set in the range from 22 to 24 °C and from 44 to 49 %, respectively. Unloaded and TrA-loaded PLGA nSMDs were obtained by spraying PLGA-EA solution at 3 % w/v. In particular, as for TrAloaded devices, TrA powder was added to the organic phase at 1:20, 1:10, 1:5 drug:polymer weight ratio. The samples were named TA5, TA10 and TA20, respectively. 2.4 Morphological analysis The obtained PLGA particles were characterised using a field emission scanning electron microscope (FE-SEM, FEI QUANTA200FEG, Netherlands) operating at 20 kV and at a working distance of 10 mm. To obtain SEM images, nSMDs were collected on a conductive metal stub and coated by a Pd–Au thin layer (about 19 nm) before FESEM analysis. Particle size and size distribution were determined by Image J analysis software—assisted (NIH) digital planimetry. The digitalised images were used to measure the area of the nSMDs, which was estimated, after image calibration, from the projected area. The diameter of the equivalent circumference, having the same projected area of the devices, was calculated and considered as an estimate of particle size. Particle circularity was obtained by ImageJ software and represents the ratio between the perimeter of the equivalent circumference and the perimeter of the actual particle. 2.5 Release kinetics Release kinetics of TrA from PLGA nSMDs were evaluated by placing the samples in PBS (120 mM NaCl, 2.7 mM KCl, 10 mM phosphate salts; pH = 7.4), in a thermostatic bath at 37 °C under agitation (100 rpm). Sink conditions were ensured in all cases. At scheduled time intervals, all the release medium was withdrawn, replaced with fresh PBS and analysed for TrA content by spectrophotometric assay (k = 241 nm). The linearity of the response was assessed in the 0.1–10 lg/mL concentration

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range (r2 [ 0.99). Results are expressed as fraction of TrA released ± SD of three repeats. Moreover, to investigate the key phenomena controlling TrA release, the released fraction was modeled by the equation proposed by Corrigan and Xue [22]. Briefly, drug release from PLGA is considered to be governed by two separate contributions. The first one takes into account the initial burst effect, i.e. the drug dissolution at the interface between PLGA and PBS, which is governed by a diffusion mechanism. The second contribution accounts for the release of the drug entrapped in the polymer, which is controlled by polymer bulk degradation. Thus, the first contribution can be described by Eq. (3): ub ¼ ub;1 ½1  expðkb tÞ

ð3Þ

where ub is the ratio between the drug amount released due to burst effect and the overall drug within the devices at time t, ub,? is the burst fraction after infinite time in the absence of other release phenomena and kb is the rate constant associated with the burst. Bulk degradation refers to (1–ub,?), i.e. the drug fraction within the polymer, and is described by the following Eq. (4), which takes into account the well-known autocatalytic degradation kinetics of PLGA [23–25]. The drug fraction released due to polymer degradation, udeg, is described by: u ¼ 1  ub;1




exp½k ð1  tmax Þ : 1 þ exp½k ð1  tmax Þ

ð4Þ

Here kdeg is the degradation constant, while tmax is the time at the maximum release rate. The overall released fraction, therefore, is given by the sum of the contributions ub and udeg: utot ¼ ub;1 ½1  expðkb tÞ
exp½k ð1  tmax Þ : þ 1  ub;1 1 þ exp½k ð1  tmax Þ

ð5Þ

Equation (5) describes drug release profile and can be solved by nonlinear least squares fitting [26] using ub,?, kdeg, kb and tmax as adjustable parameters. This equation describes drug release profiles from biodegradable devices [22, 27], and was modified for molecularly dispersed drugs in the polymers [28].

3 Results and discussion In Fig. 1 the viscosity (g) of PLGA solutions in EA versus concentration (0.01/5 % w/v) is shown in a log–log plot. Data were fit by a power law and, as displayed in the figure, the slope of the fitting functions sharply changes around 1 % w/v. In particular, g is proportional to C0.0374 below

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10

-2

10

2

10

1

10

0

G'

4% 3%

10-1 10

-2

10

-3

10-4 -1 10

2%

y = 7.72e-4 * x R= 0.970

10

0

10

1

10

G''

1.45

G', G'' (Pa)

Viscosity [Pa s]

Viscosity (Pa s)

10

10 0

2

Shear rate [1/s] -3

10

-1

10

-2

10 y = 7.37e-4 * x R= 0.816

0.0374

-3

10 -4

10

-2

10

-1

0

10

10

1

10

-4

10

C (% w/v)

Fig. 1 Viscosity as a function of PLGA concentration (0.01/5 % w/v) in EA at a shear rate value of 10 s-1. Inset: flow curves of PLGA solutions at 2, 3, 4 % w/v

1 % w/v, while g is proportional to C1.45 above that threshold value. This slope change, as reported in literature, corresponds to the overlap concentration (C*), at which the transition from the dilute to the semidilute regime occurs. Above C*, the polymer chains begin to overlap with one another and ES takes place [29]. The inset of Fig. 1 shows the rheological flow behavior of PLGA solutions at 2, 3 and 4 % w/v. A pseudoplastic behavior (shear thinning) was observed for all the tested solutions since the viscosity was found to be decreasing when increasing the shear rate. In particular, in the low shear rate region, the viscosity was sharply decreasing (thinning) while in the higher shear rate regime, the viscosity was only slightly dependent upon the shear rate. This rheological behavior is typical of macromolecular solutions and is controlled by the rate of formation/disruption of polymeric chain overlapping. Indeed, as the shear rate increases, the rate of entanglement disruption becomes predominant over the rate of entanglement formation, thus leading to the thinning. The entangled solution behavior is also evidenced by the mechanical spectra (i.e. the dependence of the elastic and viscous moduli upon frequency) of PLGA solution in EA, as shown in Fig. 2, in which the mechanical spectra of a representative solution at 3 % w/v is shown. At low frequency the solution presented a prevalently viscous behavior (G00 [ G0 ) while at high frequency a predominantly elastic behavior (G0 [ G00 ) was observed. This transition, indicated by the crossing of the G0 and G00 curves, occurred at a given value of the frequency (f*) corresponding to the intrinsic rate of disentanglement of polymer chains. In general, during ES process, polymer chain entanglements oppose the Coulomb fission, which is the process occurring when a charged droplet emits small, highly charged droplets [30]. Coulomb fission takes place when

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-1

10

10

0

10

1

2

10

Frequency (Hz)

Fig. 2 Mechanical spectra of PLGA solution in EA at 3 % w/v

the charged drop approaches the Rayleigh limit [31], which is a limiting value determined by the balance between the electric stress and surface tension. Shenoi et al. [32] identified, as a discriminating parameter defining the transition between fiber and bead formation (i.e. between electrospraying and electrospinning), the entanglement number in solution, ne,sol. The latter is defined as the ratio between polymer number average molecular weight (Mn) and Me, which is the average molecular weight of the polymer segments between two entanglements in solution. Three different regimes in the ES of polymer solutions can be thus identified: when ne,sol \ 2, regular ES (i.e. bead formation) is observed, while when ne,sol [ 3.5, electrospinning (i.e. fiber formation) appears because of the onset of a sufficiently strong elastic network that stabilises the jet against breakup. When 2 \ ne,sol \ 3.5, a mixed regime yielding beads and fibers takes place. According to the rubber-elasticity theory [33], Me can be estimated from the elastic modulus (G0 ) measurements. Briefly, taking into account that G0 is proportional to the number of entanglements, the elastic modulus can be expressed by: G0 ffi RTz

ð6Þ

where RT is the thermal energy and z the number of the entanglements expressed in mol/volume; the latter can is expressed by: c z ð7Þ Me where c is polymer concentration. Substituting in Eq.(6), Me can be immediately estimated by the following equation: Me ffi

RTc G

ð8Þ

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The ‘‘dangling ends’’, i.e. the polymer chain segments attached to the network by only one entanglement point, do not contribute to the G0 value, because they cannot store elastic energy. Therefore, Eq. (8) must be corrected as follows [34, 35]:   RTc Me Gffi 12 ð9Þ Me Mn In regard to the tested PLGA concentrations in EA, within the range of predominantly elastic behavior, in particular at 10 Hz, the obtained G’ values were in the 7–10 Pa range. Hence, being PLGA molecular weight in the 7–17 kDa range, and the calculated Me in the 3.5–8.5 kDa interval, ne was correspondingly around 2. In particular, this ne value corresponds to the condition of chain overlapping, in which the average distance between adjacent chains is of the same order of magnitude of chain length. Therefore, in our case, the sprayed solution (3 % w/v) is basically in the overlap condition. At this concentration an insufficiently deformable entangled network of polymer chains exists and thus, according to the theory, ES, and therefore beads formation, was favored [31]. Once in the ES regime, the particle morphology is determined by the competition between chain entanglement and Coulomb fission within a single droplet. More specifically, as the solvent evaporates from the droplet, two competing factors occur. The first is the increase of polymer concentration and, consequently, of entanglements which stabilise the droplet from further subdivision thus preserving its spherical shape. The second one is related to the increase of surface charge which can overcome the surface tension thus leading to droplet fission and the formation of ‘‘offspring droplets’’. In this latter case, fission tends to deform the particle into irregular shapes that may relax to the spherical one if the droplet is still in the liquid state or if the elasticity of the entangled network present in the offspring droplets is able to recover the spherical shape. In an intermediate case, in which solvent evaporation in the offspring droplets is almost complete so that the droplet may remain frozen in the shape at the time of fission, the formation of non-spherical particles is likely to be favored [31]. This latter situation is of course favored at high values of flow rates and voltage. Thus, we set experimental conditions in order to hinder the formation of spherical particles. In Fig. 3a, SEM image of unloaded nSMD is reported. As it can be seen from the figure, non-spherical microdevices were obtained, and in particular the formation of erythrocyte-like nSMDs was observed. In Figs. 3b–d, SEM images of nSMDs loaded with TrA at 1:20, 1:10 and 1:5 drug/PLGA weight ratio are reported, respectively. It can be noted that drug presence has only a slight influence on device shape and dimensions. On the contrary, the presence of the drug does alter nSMD surface properties since the drug loaded devices present a

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smoother surface. Actually, the nSMDs retained their nonspherical shape since the drug presence is not expected to influence PLGA chain entanglements. On the contrary, surface appearance is determined by the competition between solvent evaporation from droplet surface and polymer diffusion. This phenomenon is obviously associated to the polymer/solvent interaction which can be influenced by the presence of a third component in solution, i.e. the drug. Indeed, it is known that the polymer/ solvent interaction and, therefore, the solvent choice, is crucial since it affects the timescale of solvent evaporation relative to polymer diffusion in the offspring droplets. This, in turn, influences the formation of a gradient of polymer concentration between the surface and the central region of the droplet, which can trigger the formation of a glassy skin around the droplet. The properties of this external layer directly influence device architecture and morphology, accounting for the formation of erythrocyte-like or hollow particles. Moreover, phase separation induced by solvent evaporation can lead to the production of porous particles [36]. At high values of drug loading (1:5 drug/PLGA weight ratio) the presence of crystallised drug is observable (Fig. 3d), which indicates a limit in the drug loading capability at this values of voltage and flow rate. The size of nSMDs was assessed by ImageJ software, as described in Sect. 2.3. In particular, in all the ES conditions, nSMD diameters of the projected area ranged between approximately 4 and 11 lm, without a clear trend with applied voltage, tip distance and flow rate (Fig. 4). Likewise, the circularity of nSMD projected area, as calculated by ImageJ software, did vary between 0.65 and 0.85 irrespective of drug loading and ES parameters (Fig. 5). This results suggest that, once in the mixed regime of non-spherical particle formations, a good reproducibility can be attained being the size of the devices slightly influenced by process parameters. TrA release profiles are reported in Fig. 6. Basically, the same fractions are delivered in approximately 3 weeks and no significant burst was observed at any drug loading. Moderate acceleration of release rate was envisaged in the case of high loading formulation (1:5 drug:PLGA weight ratio), which suggests significant fractions of superficial drug in this case. Released fraction was modeled by the equation proposed by Corrigan and Xue [22] and the results are summarised in Table 1. The burst constant kb showed relatively low values, which is consistent with the very low water solubility of TrA. Moreover, ub,? was decreasing from about 0.3, for TrA:PLGA ratio of 1:5 and 1:10, to 0.214 when TrA:PLGA ratio is 1:20. In this latter case, most of the drug is embedded within the polymeric matrix, while in the case of 1:5 ratio, a significant amount of superficial TrA is present, and is quickly released, as shown in release profiles of Fig. 6. The

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Fig. 3 SEM images of PLGA nSMDs. a unloaded; b TA5; c TA10; d TA 20 devices. The bar is 5 lm

16 Unloaded TA 1:20

14

TA 1:10

12

TA 1:20

d, µm

10 8 6 4 2

3 0. 41 18 -

18 -0 .2 -1 3

0 0. 41 18 -

18 -0 .2 -1 0

3 0. 41 15 -

15 -0 .2 -1 3

0 0. 41 15 -

15 -0 .2 -1 0

0

Fig. 4 MD diameters of the projected area of unloaded and TrA—loaded MDs with TrA at 1:20, 1:10 and 1:5 drug/PLGA weight ratio at different applied voltage, tip distance and flow rate

degradation constant kdeg was weakly affected by drug loading, being similar for 1:5 and 1:10 formulations suggesting that degradation occurs by the same pattern in these cases. Indeed, as shown in SEM image (Fig. 3d), in the case of 1:5 mass ratio, superficial TrA crystals are present,

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and this suggests that a saturation limit of drug is reached in nSMD. In the case of 1:20 TrA:PLGA ratio a higher value of the degradation constant kdeg was found and this can be explained considering that the autocatalytic degradation of the PLGA is enhanced when a lower percentage

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1

Circularity

0.8

0.6

0.4 Unloaded

0.2

TA 1:20 TA 1:10 TA 1:5

13 -0 .4 -

18 -0 .2

18

-1 3

0 18

-0

.4

-1

0 18

15 -0 .4

-0 .2 -1

-1 3

3 -0 .2 -1 15

15 -0 .4

15

-0

.2

-1

-1 0

0

0

Fig. 5 Circularity results of unloaded and TrA—loaded MDs with TrA at 1:20, 1:10 and 1:5 drug/PLGA weight ratio at different applied voltage, tip distance and flow rate

Released TrA fraction

1

Table 1 Results of modeling release kinetics of MDs loaded with TrA at 1:20, 1:10 and 1:5 drug/PLGA weight ratio

0,8

TrA:PLGA 1:5 w/w

TrA:PLGA 1:10 w/w

TrA:PLGA 1:20 w/w

kb (day-1)

0.890 ± 0.184

0.366 ± 0.266

0.699 ± 0.32

ub,?

0.299 ± 0.037

0.301 ± 0.148

0.214 ± 0.149

kdeg (day-1)

0.338 ± 0.009

0.332 ± 0.017

0.393 ± 0.014

10.6 ± 0.3

9.69 ± 1.11

8.57 ± 0.36

0,6 0,4 0,2 TA 1:20 0 0

5

10

15

20

tmax (day)

25

t, day Released TrA fraction

1

of drug is present within the polymeric matrix. Moreover, the trend of tmax with drug loading was consistent with the one of the degradation constant. This parameter was, indeed, about 10 days for the 1:5 and 1:10 formulations, while being lower for 1:20 formulation. This occurs because higher degradation rates are associated to shorter times needed to get the maximum degradation rate.

0,8 0,6 0,4 0,2 TA 1:10 0 0

5

10

15

20

25

t, day

4 Conclusions

Released TrA fraction

1 0.8 0.6 0.4 0.2 TA 1:5 0 0

5

10

15

20

25

t, day

Fig. 6 Release profiles in PBS from MDs loaded with TrA at 1:20, 1:10 and 1:5 drug/PLGA weight ratio

Results of this study showed that the ES of polymer solutions with suitable viscosity allows the fabrication of nSMDs made up of PLGA. The addition of TrA did not affect polymer chain interactions and, as a consequence, the particle morphology, independently on the drug loading. This is particularly appealing in the pharmaceutical industry for the ease of scaling-up of this technique. Furthermore, the drug could be released for a prolonged period, with controlled and reproducible kinetics for over 3 weeks. Finally, modeling results highlighted that release is mainly driven by degradation, at a higher extent in the case of low drug loading.

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390 Acknowledgments This study was financially supported by the National Operative Programme REPAIR (PON01-02342). Scanning Electron Microscopy was supported by the Transmission and Scanning Electron Microscopy Labs (LAMEST) of the National Research Council.

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