Film coatings for oral pulsatile release

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ARTICLE IN PRESS International Journal of Pharmaceutics xxx (2013) xxx–xxx

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Review

Film coatings for oral pulsatile release Alessandra Maroni, Lucia Zema, Giulia Loreti, Luca Palugan, Andrea Gazzaniga ∗ Università degli Studi di Milano, Dipartimento di Scienze Farmaceutiche, Sezione di Tecnologia e Legislazione Farmaceutiche “Maria Edvige Sangalli”, Via G. Colombo 71, 20133 Milan, Italy

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Article history: Received 19 January 2013 Received in revised form 6 March 2013 Accepted 10 March 2013 Available online xxx Keywords: Oral drug delivery Coated dosage forms Pulsatile release Lag time Chronotherapy Film-coating

a b s t r a c t Pulsatile delivery is generally intended as a release of the active ingredient that is delayed for a programmable period of time to meet particular chronotherapeutic needs and, in the case of oral administration, also target distal intestinal regions, such as the colon. Most oral pulsatile delivery platforms consist in coated formulations wherein the applied polymer serves as the release-controlling agent. When exposed to aqueous media, the coating initially performs as a protective barrier and, subsequently, undergoes a timely failure based on diverse mechanisms depending on its physico-chemical and formulation characteristics. Indeed, it may be ruptured because of the gradual expansion of the core, swell and/or erode due to the glassy-rubbery polymer transition or become permeable thus allowing the drug molecules to diffuse outwards. Otherwise, when the coating is a semipermeable membrane provided with one or more orifices, the drug is released through the latter as a result of an osmotic water influx. The vast majority of pulsatile delivery systems described so far have been prepared by spray-coating, which offers important versatility and feasibility advantages over other techniques such as press- and dip-coating. In the present article, the design, manufacturing and performance of spray-coated pulsatile delivery platforms is thus reviewed. © 2013 Elsevier B.V. All rights reserved.

Contents 1. 2. 3. 4. 5. 6.

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Rupturable film coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Erodible film coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Permeable film coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Semipermeable film coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

1. Introduction Pulsatile delivery systems are non-conventional dosage forms designed to release the active ingredient after a lag phase of programmable duration thereby allowing a chronotherapeutic effect to be attained (Maroni et al., 2005, 2010). Most current pulsatile delivery systems are typically time-controlled in that the onset of release is prompted by inherent mechanisms irrespective of the differing conditions they may encounter in the outer environment. Among them, formulations intended for the oral route are of particular interest in the case of chronic pathologies with circadian symptoms that have a high likelihood of recurring in the night

∗ Corresponding author. E-mail address: [email protected] (A. Gazzaniga).

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or early morning hours, such as cardiovascular disease, bronchial asthma, rheumatoid arthritis and sleep disorders. Indeed, medications able to provide an appropriate delay phase prior to drug release, administered at bedtime, could selectively cover the especially critical period during which the disease state tends to worsen with no need for waking up the patient for drug intake. Apart from chronopharmaceutical purposes, oral time-based systems for pulsatile release can be exploited to achieve colon delivery, provided that the dosage form is enteric coated, to overcome unpredictable gastric residence, and the in vivo lag phase is at least as long as the relatively consistent small intestinal transit time (SITT, 3 ± 1 h, mean ± SE) (Davis, 1985; Gazzaniga et al., 1994a, 2006). In the oral delivery area, colonic release is a major focus of research mainly because of the established benefits it may offer in the therapy of inflammatory bowel disease (IBD) (Friend, 2005) and of its potential as a release site for peptide drugs, which generally

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Fig. 1. Outline of the performance of coated delivery systems for oral pulsatile release on exposure to aqueous fluids.

show oral bioavailability issues (Haupt and Rubinstein, 2002; Bourgeois et al., 2005; Maroni et al., 2012). In addition, oral pulsatile delivery devices, particularly when able to yield multi-pulse release profiles, may serve in place of prolonged-release systems with drugs that are subject to a strong first-pass metabolism or develop pharmacological tolerance (Bussemer et al., 2001; Stubbe et al., 2004) and, in the specific case of antibiotics, would limit the growth of resistant bacterial strains by circumventing defensive dormancy and affecting a larger number of micro-organisms in the division phase (Saigal et al., 2009). Moreover, it has recently been suggested that the use of pulsatile release dosage forms could prevent detrimental interactions between co-administered drugs from occurring within the gastrointestinal (GI) tract (Sawada et al., 2003). Although the earliest pulsatile delivery formulations were devised as multi-layer tablets partially enclosed in an impermeable shell (Conte et al., 1989, 1992), a timed liberation of orally-administered bioactive compounds is currently achieved mainly by the application of a functional polymeric coating to a drug-containing core. The core may either be a single- or a multiple-unit dosage form, the latter enabling improved reproducibility in the GI transit and absorption consistency (Bianchini et al., 1992, 1993). The performance of the coating strictly depends on the relevant physico-chemical nature and is started on exposure to the aqueous biological fluids (solvent activation). Accordingly, rupturable, erodible, permeable and semipermeable layers can be distinguished (Fig. 1) (Bussemer et al., 2001; Maroni et al., 2005, 2010). In this respect, various coating techniques, such as spraycoating, double-compression, dipping and powder-layering, have been employed according to the type of materials involved and the features desired for the final coat. However, spray-coating, i.e. the process whereby a solid substrate is provided with a thin layer of polymeric material deposited from a nebulized suspension or solution of the polymer itself, is by far the most widely used owing to its advantageousness in terms of (i) time, costs and scalability of the process, (ii) homogeneity of the thickness, structure and surface of the coat, (iii) fine modulation of the coating level and (iv) versatility with respect to the type and dimensions of the starting cores. Therefore, its applications in this particular field are herein reviewed. 2. Rupturable film coatings Rupturable coatings are insoluble though moderately permeable polymeric films that undergo timed disruption following

exposure to the aqueous media thereby allowing the drug to be released after a programmed lag phase. Their disruption is induced by the hydrostatic pressure that builds up inside the core mainly as a result of the swelling of hydrophilic polymers or of an osmotic water influx. The extent to which drug release is delayed generally depends upon the thickness and composition of the rupturable film. Ethyl cellulose (EC), often in admixture with plasticizers and/or channelling excipients that improve its inherent flexibility and permeability characteristics, has most frequently been employed as the film-forming agent. In a number of instances, EC films are superimposed on thicker layers formed from highly waterswellable polymers, such as the superdisintegrants croscarmellose sodium (Ac-Di-Sol® ) and low-substituted hydroxypropylcellulose (L-HPC), that provide, once in a hydrated state, the outward pressure required for rupture. The first example of formulations with such design features was the Time-Controlled Explosion System (TES), which consisted of sucrose seeds of 350–500 ␮m in diameter overlaid with drug, L-HPC and EC layers, respectively (Ueda et al., 1994a,b,c). Thus, the manufacturing of TES involved three different steps: the drug and the swelling polymer were successively loaded onto nonpareils by powder-layering in a centrifugal granulator using a low-viscosity hydroxypropylmethylcellulose (HPMC) dispersion as the binder, whereas EC was applied by spraying a 3:2 ethanol/dichloromethane solution of the polymer in a fluid bed. After each coating step, the units underwent 24 h drying in a vacuum oven at 40 ◦ C. While the thickness of the EC film affected the delay time, as demonstrated by testing values in the 23–43 ␮m range, that of the swelling L-HPC layer was critical to the achievement of a pulsatile release behaviour, 180 ␮m being the threshold beyond which such a performance was observed. Differing TES formulations showed consistent in vitro release data irrespective of the dose and solubility of the active ingredient, size of the subunits and pH of the dissolution fluid. The addition of talc to the EC membrane resulted in a decreased mechanical strength of the latter and consequently shortened the delay phases (Ueda et al., 1994c). Evaluated in dog and human pharmacokinetic studies, TES provided in vivo lag times in agreement with the in vitro ones (Ueda et al., 1994d; Hata et al., 1994; Murata et al., 1998). When systems having longer in vitro lag times were administered to dogs, a decrease in bioavailability was found, which was ascribed to the limited water content of distal intestinal fluids or to an enhanced first-pass effect possibly connected with a slower absorption process. In the case of prototypes with a shorter in vitro lag time (3 h), the plasma concentration profiles of the vasodilator drug employed revealed no major influence of the fed/fasted state of the volunteers (Fig. 2) (Hata et al., 1994).

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Fig. 2. Mean blood concentration profiles of FK409 from immediate-release formulations in the fed state () and TES delivery systems in the fasted (䊉) as well as fed () state (9 volunteers; bars indicate standard deviation). Reproduced with permission from Hata et al. (1994).

Rupturable EC films and swelling polymeric layers were also combined within single-unit formulations, such as in a delivery system prepared from hard- (size 0 or 3) and soft- (12.6 mm × 8 mm) gelatin capsule cores (Bussemer et al., 2003a,b; Bussemer and Bodmeier, 2003; Dashevsky et al., 2004). The choice of Ac-Di-Sol® and EC as the expanding and rupturable agent, respectively, was based on preliminary testing of free cast films composed of differing polymeric materials with potentially suitable characteristics (Bussemer et al., 2003b,c). Both the EC and Ac-Di-Sol® layers were prepared by organic spray-coating carried out in a lab-scale pan coater. Ac-Di-Sol® suspensions, in either isopropanol or a 96:4 (v/v) ethanol/water blend, contained polyvinylpyrrolidone (PVP, Kollidon® 30 or 90 F) as a thickener preventing sedimentation of the polymer particles within the coating operations. PVP also served as a binder promoting the formation of a continuous swelling layer and its adherence to the core surface. Moreover, it could aid the disintegration of such a layer following hydration, so that this would not hamper a fast release of the drug from the delivery system at the end of the delay period. On the other hand, EC was dissolved in ethanol/water mixtures at 90:10 or 96:4 (v/v), with or without dibutyl sebacate (DBS) as the plasticizer. Further additives, such as HPMC and magnesium stearate, could be included in the EC coating solution. Analogous operating conditions were adopted for the two coating steps. The product was consistently kept at mild temperatures (<30 ◦ C) reasonably because of the heat-sensitive sticking tendency of the gelatin substrate. Post-coating drying was performed inside the pan for relatively short time lapses under temperatures comparable to those set for the process. In some cases, the coated units were subjected to additional drying in oven. Coating levels in the order of tens and units of mg/cm2 were reached with Ac-Di-Sol® and EC, respectively. The obtained systems showed a typical pulsatile release behaviour, and the lag phase was prolonged by either increasing the thickness of the EC film or reducing the swelling polymer coating level. Longer lag phases were observed with hard-gelatin capsule cores, probably due to a partial discharge of the swelling pressure developed by the expanding layer towards the inner powder-filled shell, which would not be the case with liquid-filled soft-gelatin units. The batch-to-batch coating level uniformity was improved by

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increasing the distance separating the spray nozzle from the substrate and the pan rotational speed, provided that a threshold of 30 rpm was not exceeded to avoid circle-like motion patterns, or alternatively underloading the process chamber (Dashevsky et al., 2004). Moreover, the use of standard baffles turned out preferable. Only when tablets were coated in place of capsule cores, an increase in the amount of Ac-Di-Sol® applied resulted in extended delay phases (Sungthongjeen et al., 2004). This was ascribed to the barrier function fulfilled by the swollen polymer that would hinder the tablet disintegration process. The addition of magnesium stearate weakened the EC membrane by lowering its tensile strength. A shorter time was therefore necessary for film disruption, which was also obtained by raising the amount of pore-former (low-viscosity HPMC) or reducing that of plasticizer (Bussemer and Bodmeier, 2003; Sungthongjeen et al., 2004). Such formulation changes were indeed related to a faster penetration of water into the swelling layer and a lower flexibility of the rupturable film, respectively. In order to limit the use of organic solvents, the EC membrane was subsequently applied by aqueous spray-coating (Mohamad and Dashevsky, 2006a). An EC dispersion (Aquacoat® ECD) diluted to a 15% (w/w) solid content and plasticized with triethyl citrate (TEC) was employed. In this case, more plasticizer had to be added. Besides, the spray rate was diminished to avoid pre-swelling of the underlying Ac-Di-Sol® layer, and stronger operating as well as curing temperature conditions were needed for solvent removal as compared with organic spray-coating. However, the films prepared with aqueous EC dispersions were less effective in delaying drug release, and considerably higher coating levels were thus required. The issue of the pH-dependent swelling of Ac-Di-Sol® was overcome through the incorporation of fumaric acid, which could maintain an acidic microenvironment inside the hydrated polymer matrix thereby enabling a higher reproducibility of its performance within the rupturable pulsatile delivery system. According to the same design concept based on coupled Ac-Di-Sol® and EC coatings, pellet formulations were prepared starting from sugar spheres of 355–500 ␮m (600 g batch size) by successively layering ethanolic solutions or suspensions of the drug and the swelling polymer, containing plasticizing and binding excipients, and, finally, a TEC- or DBS-plasticized 15% (w/w) aqueous dispersion of EC (Mohamad and Dashevsky, 2006b, 2007). The product temperature was kept at 26 ◦ C or less, and drying was carried out for 15 min at 40 ◦ C. As in the analogous single-unit preparations, the outer membrane disruption was confirmed to be necessary for the drug to be released. Delay times were longer in the presence of a strongly acidic pH or of Na+ and Cl− ions as compared with purified water (Mohamad and Dashevsky, 2007). However, they were unaffected by the hydrodynamic conditions. The administration of paracetamol pellets with increasing EC coating levels to healthy volunteers resulted in a progressive decrease in the extent of absorption vs. an immediate-release marketed product. It was concluded that the observed loss of bioavailability would not depend on the performance of the delivery system itself but rather on the limited water content of colonic fluid, which might impair a quantitative dissolution for highly-dosed drugs with moderate solubility. In order to prepare an analogous pellet formulation, starting cores in the 20–24 mesh size range containing isosorbide 5-mononitrate, the main active metabolite of isosorbide dinitrate, were manufactured by extrusion-spheronization with lactose and microcrystalline cellulose following statistical optimization of the relevant composition and process parameters (Liu et al., 2009). The swellable and the rupturable polymer were applied in a fluid bed by aqueous spray-coating, nebulizing a solution of croscarmellose sodium and a dispersion of EC (both at 15%, w/w) under the same operating conditions, the inlet air temperature being adjusted to 38 ◦ C. At the end of each coating step, drying was continued at that temperature for 15 min and 2 h, respectively. In view of the

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programmable delay that was consistently observed in vitro and in beagle dogs prior to release of the anti-anginal drug from pellets with 20% croscarmellose sodium and 16% EC weight gains, a possible chronotherapeutic use of this delivery system for the management of nocturnal cardiovascular events was suggested in place of prolonged-release formulations that would worsen the tolerance issues associated with the concerned active ingredient. In another multiple-unit system based on layered pellets, low- and high- viscosity grades of HPMC were employed as swelling agents instead of croscarmellose sodium (Yadav et al., 2011). The core pellets had a diameter of 600–900 ␮m and were prepared by extrusionspheronization from microcrystalline cellulose, lactose, sodium laurylsulphate and the second-generation sulphonylurea glipizide. Spray-coating in fluid bed was first performed with 5% w/v water solutions of HPMC and afterwards with a 5% w/v ethanolic solution of EC containing TEC (25% on polymer) on small product batches (60 g). The inlet air temperature was higher (∼70 ◦ C) during the HPMC coating process. Two formulations, one with 20% and 5% weight gain of low-viscosity HPMC and EC, respectively, and the other with 5% and 10% weight gain of high-viscosity HPMC and EC, were finally selected for combination with uncoated pellets to provide double-pulse delivery patterns with a lag phase of 6–8 h between the former and the latter pulse. Such patterns were expected to be advantageous in terms of reduced dosing frequency and, consequently, enhanced patient compliance with the chronic hypoglycaemic treatment. Besides being applied as individual layers underlying the rupturable EC film, swelling excipients were also mixed with the components of the core formulation. In the Swelling Controlled Release System (SCRS), films composed of TEC (20% on polymer)-plasticized EC and low-viscosity HPMC as a pore-former surrounded a 5 mm core tablet including polyvinyl alcohol (PVA) as the swellable agent (Morita et al., 2000a). Because of their swelling-limiting properties, inorganic or carboxylic acid salts, such as trisodium citrate dihydrate, were optionally added. The rupturable film was applied by spray-coating using a 5% hydroethanolic (20:80, v/v) solution of EC. By modifying the EC/HPMC and PVA/swelling-limiting salt ratios, it was possible to modulate the release of the model drug emedastine difumarate as regards the onset and the rate of release. In particular, formulations having an EC/HPMC ratio of 75:25, a high PVA content and no swellinglimiting salts showed a lag phase prior to a rapid liberation of the active ingredient, which coincided with the EC film breakup and could be delayed as a function of the coating level. SCRS systems with zero-order release after the lag phase were studied in healthy volunteers providing plasma concentration data in agreement with in vitro release ones when a purposely adapted flow-through dissolution testing method, with improved biorelevance in terms of fluid volume, was employed (Morita et al., 2000b, 2003). In an analogous delivery system, differing viscosity grades of HPMC were incorporated in biconvex tablets of 6 mm that were coated (6 kg batches) with hydro-ethanolic solutions (95%, v/v ethanol) of EC containing high-viscosity HPMC as the channelling agent (Lin et al., 2008). In media with diverse pH, ionic strength and osmolality characteristics, the resulting coated systems delayed the liberation of the vasodilator doxazosin mesilate as a function of the thickness and HPMC content of the outer membrane. The use of coating solutions having a larger volume of water was demonstrated to shorten the lag phase. On the other hand, the release rate was dependent on the viscosity grade of the swellable polymer included in the core. The plasma concentration curves obtained following administration to healthy volunteers of prototypes with 2.7 to 8% weight gains and 1:1 to 1.8:1 EC/HPMC coating ratios were in line with the in vitro results, pointing out a higher degree of correlation with lag time data from pH 1.2 and release rates from pH 6.8 fluids. Alternatively, the

superdisintegrant crospovidone was employed as the swelling agent for a multiple-unit formulation wherein the cores were pellets with differing amounts of sodium chloride, manufactured by various techniques (wet high-shear pelletization, melt pelletization and prilling), whereas the rupturable film consisted of plasticized EC applied from a 30% w/w aqueous dispersion with 15% w/w on polymer of DBS (Hartman Kok et al., 2001). The coated pellets were dried for 12 h in an oven at 60 ◦ C. Pulsatile release profiles with lag times <1 h, as assessed by conductivity measurements, were achieved from systems with 10–27 ␮m thick EC films. A faster release process was observed when dealing with core pellets that exhibited a uniform shape and narrow size distribution range, i.e. the ones prepared by spraying a molten PEG 4000based mixture and allowing the resulting droplets to solidify in cool paraffin oil (prilling). This was reflected, especially in the case of the higher coating levels, in a lesser inter-pellet variability of individual release profiles and, ultimately, in a more pronounced slope of the cumulative curves. A Eudragit® L sub-coating was optionally applied with the aim of smoothing the especially rough surface of pellets prepared by wet high-shear pelletization, thus enhancing the quality of superimposed EC films. However, the presence of a pH-dependent interlayer might deeply alter the overall release mechanism of double-coated pellets and impair the reproducibility of their performance depending on the GI site of break-up. As compared with formulations without crospovidone, those provided with the swelling agent showed shorter lag phases but unchanged release rates, thereby indicating that, with respect to the osmotic agent alone, such an excipient could bring about an earlier though not a more complete disruption of the outer film. Crospovidone was also contained in a tablet core (7 mm) that was coated with ethanolic mixtures of EC and Eudragit® L (1:2) by a pan coater (Fan et al., 2001). The lag time before liberation of the calcium channel blocker diltiazem hydrochloride could be modulated by varying the coating level from 6 to 8.4% without impacting on the release rate. The performance observed in vitro was confirmed by a pharmacokinetic study in 8 volunteers, which highlighted an agreement between in vivo absorption and in vitro release data as well as the absence of significant differences in terms of AUC0–24 vs. a conventional tablet preparation. In another system, EC was mixed with Eudragit® L and applied, by hydro-alcoholic (95%, v/v ethanol) pan coating, to a hydrophilic matrix of 8 mm in diameter formed from high- and low-viscosity HPMC (Feng et al., 2008). A weight gain of 10% was approximately reached. Such a device exhibited a delayed-onset prolonged release of propranolol hydrochloride both in vitro and in beagle dogs. In the above-described formulations, Eudragit® L was expected to act as a pH-dependent pore-former ensuring gastric resistance properties (Fan et al., 2001; Feng et al., 2008). After its dissolution at the enteric pH, the hydrophilic polymers present in the tablet and matrix cores, i.e. crospovidone and HPMC, would indeed be allowed to swell due to the ingress of water and thus promote a complete break-up of the insoluble discontinuous film. However, a timely onset of drug release might fail to be achieved because of the unpredictability of gastric residence, and this would negatively reflect on the outcome of cardiovascular disease chronotherapy. In a particular instance, the hydrostatic pressure required for breakage of the EC film was exerted by carbon dioxide, which was generated inside a biconvex core tablet (9 mm) following solubilization of the effervescent excipients it contained along with chlorpheniramine maleate (Krögel and Bodmeier, 1999). The rupturable membrane was applied by spraying, in a lab-scale fluid-bed apparatus, a 10% w/v ethanolic solution of EC with 20% on polymer of DBS. The in vitro delay phase was extended by increasing the coating level or the core hardness. When the more flexible and water permeable film-forming polymer Eudragit® RL was utilized

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PEG 6000 were added to the Eudragit® RS/Eudragit® RL mixtures along with talc as an anti-tacking agent. At the end of the coating processes, solvent residues were removed through a drying phase of 5 h at 50 ◦ C, which was higher than the product temperature maintained within the two coating steps. Indeed, a maximum of 38 ◦ C was reached during the HPMC coating. Upon exposure to the aqueous medium, the outer film was subject to the formation of micrometric fissures due to the mechanical stress undergone following water influx and resulting expansion of the core. The drug was thus released by diffusion and osmotic pumping after a lag time that was dependent on the coating level (1.5–10%) regardless of differing pH and hydrodynamic testing conditions. As demonstrated by the use of release media with increasing osmolality values, the osmotic mechanism was prevalent. The swelling and the osmotic excipient were both needed in order to attain a fast release of the drug after the delay period. Fig. 3. Mean release profiles of paracetamol from pellets coated with cellulose acetate up to 6 mg/cm2 (,♦) and 10 mg/cm2 (,), with (,) or without (♦,) sodium chloride in the core (paddle dissolution apparatus, 900 ml of pH 7.4 phosphate buffer, 37 ◦ C, 100 rpm, 2 replicates; bars indicate standard deviation). Reproduced with permission from Schultz and Kleinebudde (1997).

instead of EC, a floating formulation was obtained because of a faster water penetration and subsequent gas entrapment into the dosage form. Apart from EC, other insoluble polymers were exploited for the manufacturing of pulsatile delivery systems provided with rupturable coatings. For example, cellulose acetate, a cellulose derivative suitable for attaining semipermeable membranes, was applied to extruded-spheronized pellets containing microcrystalline cellulose as the extrusion aid and paracetamol as the model drug (Schultz and Kleinebudde, 1997). 3 kg batches of pellets in the 710–1400 ␮m sieve fraction were coated in fluid bed with a 90:10 acetone/2propanol co-solvent mixture including TEC and talc as the plasticizing and the anti-tacking agent, respectively, at 40% and 100% on the dry cellulose acetate weight. A timed break-up of the outer film and consequent pulsatile release of the active ingredient were brought about by the presence of sodium chloride within the pellet cores at a sufficient percentage to drive a massive entry of water (Fig. 3). When the cellulose acetate coat was stretched by the expanding core, micrometric cracks were indeed formed through which the drug solution could slowly be released. By exploring the performance of systems with coating levels up to 12 mg/cm2 , it was assessed that a minimum value of 2 mg/cm2 was needed for pulsatile release to occur, and that lag time was linearly dependent on the amount of polymer applied. The addition of plasticizers with increasing lipophilicity to the coating formula resulted in longer delays and lower release rates (Schultz et al., 1997). This was attributed to a decrease in the permeability of the membrane rather than to a modification of its tensile properties, as the latter were more affected by the quantity than by the type of plasticizer incorporated. On the other hand, the duration of the lag phase was reduced by raising the amount of talc in the film, while a minor influence on the rate of release was observed. Rupturable films composed of insoluble polymethacrylates were also described. In particular, mixtures of Eudragit® RS and Eudragit® RL were applied to terbutaline sulphate-containing tablets provided with an osmotic charge (sodium chloride at 30%) and a swelling layer made of low-viscosity HPMC (Zhang et al., 2003). The HPMC and acrylic layers were both prepared by hydroorganic pan coating, using water/ethanol blends as vehicles for the coating polymers and 50 g of cores as the substrate. PEG 400 was employed as the plasticizer for the inner coat, whereas TEC and

3. Erodible film coatings Erodible coatings are mostly based on hydrophilic polymers with pH-independent solubility that, when exposed to the aqueous media, undergo swelling, dissolution and/or erosion thereby delaying the release of the active ingredient from the inner formulation (Gazzaniga et al., 2008). The duration of the lag phase is dictated by the physico-chemical properties of the employed polymeric material, above all molecular weight, and by the relevant coating level. Hydrophilic cellulose ethers, such as HPMC, hydroxypropylcellulose (HPC) and hydroxyethylcellulose (HEC), are most frequently chosen as release-delaying agents because they offer major advantages such as a well-established safety profile, swelling properties in a range of modes and degrees, ease of handling and reasonable costs. The coating technique also affects the performance of the polymeric barrier applied. The manufacturing of pulsatile delivery systems with erodible coats has mainly been accomplished by double-compression, thus leading to rather thick porous layers in place of thinner continuous films. However, the latter technique involves a number of technical drawbacks, such as the difficult centring of the core, prepared beforehand, inside the die of the tableting machine, which may impair the coat thickness homogeneity and, consequently, the achievement of reproducible lag times, and the need for special equipment as well as time-consuming multi-step processes (Gazzaniga et al., 1994c; Ozeki et al., 2004). Moreover, double-compression suffers from versatility constraints associated with the relatively large amounts of coating polymer required. Therefore, its use would be confined to single-unit formulations of small size, so that swallowing would still be ensured after application of the coating layer, and may not allow pulsatile delivery patterns with short delay and/or steep release phases to be obtained. Spray-coating was alternatively attempted in view of its marked flexibility and easier industrial scalability, although limited experience was available on the utilization of swellable hydrophilic polymers (Maffione et al., 1993; Gazzaniga et al., 1994b,c). Only the lowest viscosity grades were indeed exploited to form relatively thin films intended for sealing, taste masking or cosmetic purposes. However, high-viscosity polymers appeared more promising as release-delaying agents. Hydro-organic suspensions of highviscosity HPMC (Methocel® K4 M and K15 M) were initially used in order to overcome the typical thickening effect that such polymers exhibit when interacting with aqueous media. In particular, water/ethanol mixtures were prepared wherein the amount of ethanol was maintained above 80% w/w to improve the feasibility of the coating operations by allowing dispersions with polymer content of approximately 5% w/w to be sprayed at reasonable rates. On the other hand, the amount of water in the dispersions was not restrained below 5% w/w because, following solvent evaporation, the coalescence of the polymeric particles would be enhanced in

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Fig. 4. Individual release profiles of paracetamol from uncoated tablet cores and delivery systems coated with Methocel® E5, Methocel® E50 and Methocel® K4 M up to a 20% weight gain (modified disintegration apparatus, 900 ml of distilled water, 37 ± 0.5 ◦ C, 31 cycles/min). Reproduced with permission from Sangalli et al. (2004).

the presence of an adequate degree of hydration. The addition of auxiliary substances such as PEG 400, talc and PVP as the plasticizing, anti-tacking and binding agent, respectively, was investigated. These dispersions were applied to 250–500 g batches of biconvex tablets (4.5 mm) and minitablets (2.5 mm) by a rotating pan of 15 l capacity. After coating, a drying phase of 20 min was carried out under the operating temperature conditions (30 ◦ C for the substrate and ∼50 ◦ C for the inlet air flow). The coated units showed satisfactory physico-technological properties (homogeneous thickness and smooth surface of the coating) and yielded the pursued release performance, with reproducible lag times depending on the viscosity grade of the polymer and the coating level. However, in view of the environmental and safety-related issues raised by the use of organic solvents, the viability of aqueous spray-coating was explored with differing viscosity grades of HPMC, namely Methocel® E5, E50 and K4 M (Gazzaniga et al., 1995; Sangalli et al., 2001, 2004; Zema et al., 2007). Their water solutions were prepared 24 h beforehand, stored at 4 ◦ C to aid complete polymer dissolution and heated up to 40 ◦ C prior to use. The coating of tablets (6 or 6.8 mm in diameter, 1 or 1.5 kg batches) by a fluid bed apparatus required a progressive adjustment of the operating conditions that was chiefly intended to ensure proper “sprayability” characteristics of the polymeric solutions, i.e. the possibility of undergoing nebulization at a reasonable rate without major problems of nozzle clogging or of powdering inside the fluid bed chamber. As expected, the process time was related to the viscosity grade of the polymer. Indeed, the HPMC content of the coating solutions needed to be reduced as a function of viscosity, thus involving longer spraying and drying phases. With all the polymers under investigation, process yields above 70% were obtained, and coated systems with satisfactory physico-technological properties could be prepared. Release tests, also performed in a modified disintegration apparatus to prevent the adhesion of the swollen HPMC layer to the inner surface of the vessels, showed that the functional coatings applied were capable of delaying the drug liberation depending on the coating level and the polymer viscosity characteristics (Fig. 4). Only in the case of Methocel® K4M-coated systems, a limited percentage (<5%) of the model drug was found to slowly leach out prior to its quantitative release. Due to the inherent high viscosity properties, Methocel® K4 M proved to form a persistent gel barrier that was able to withstand extensive erosion phenomena even after complete hydration, thus allowing drug diffusion to take place until break-up was aided by the swelling force of the disintegrating core tablet (Zema et al., 2007). Because process and performance advantages were generally demonstrated to arise from the use of lowand high-viscosity HPMC grades, respectively, special parameters

Fig. 5. Mean saliva concentration profiles of antipyrine from uncoated tablet cores (F) and delivery systems coated with Methocel® E50 up to 25% (F25), 50% (F50) and 100% (F100) weight gains (4 fasting volunteers; bars indicate standard deviation). Reproduced with permission from Sangalli et al. (2001).

were introduced for a comprehensive evaluation of the differing coating agents (Sangalli et al., 2004). In particular, the Time Equivalent Process Parameter (TEPP) was the process time (min) vs. in vitro lag time (min) ratio, whereas the Time Equivalent Thickness Parameter (TETP) was the coat thickness (␮m) vs. in vitro lag time (min) ratio. TEPP and TETP would therefore indicate the process time and coat thickness required to attain a unit of lag time, respectively. In spite of favourable TEPP and TETP results, a fine modulation of lag time as a function of the coating level would hardly be possible with Methocel® K4 M in view of its marked release-delaying ability, as highlighted by the low TETP value. On the other hand, Methocel® E50 proved to offer an acceptable balance of key aspects such as manufacturing feasibility, effectiveness in deferring the drug liberation, flexibility in the delay modulation and lack of impact on the release rate (Gazzaniga et al., 1995; Sangalli et al., 2004). Moreover, the coating process based on this polymer was found reproducible, robust and potentially scalable. Indeed, a consistent performance was achieved from systems prepared with coating solutions at differing concentrations, belonging to differing batches or obtained when shifting from fluid bed to rotating pan apparatus and raising the batch size up to 15 kg. Following administration of Methocel® E50-coated formulations to fasting healthy volunteers, the lag times preceding appearance of the marker drug antipyrine in saliva correlated with the coat thickness (Fig. 5) (Sangalli et al., 2001). In addition, they were comparable with the corresponding in vitro data. Imaging studies also pointed out the ability of such systems to provide time-controlled colon delivery when in a gastroresistant configuration (Sangalli et al., 2009). The same coating procedure based on an 8% w/v Methocel® E50 solution with 0.8% w/v of PEG 400 was subsequently employed for the preparation of bovine insulin delivery systems, with or without enzyme inhibitor and absorption enhancer compounds in the core tablet (Maroni et al., 2009; Del Curto et al., 2009). Such systems exhibited the desired stability and in vitro release properties thereby demonstrating that the overall manufacturing operations, and particularly the challenging film-coating step, would not impair insulin integrity. Because an earlier liberation of the protease inhibitor and/or permeation enhancer was hypothesized to increase the intestinal absorption of proteins, “two-pulse” delivery devices were later proposed for a differential modulation of the drug and adjuvant

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release (Del Curto et al., 2011). These consisted of an insulin tablet and three coatings, i.e. a camostat mesilate or sodium glycocholate film enclosed between an inner and an outer Methocel® E50 layer. The outer coat was meant to delay the onset of the adjuvant liberation throughout the small bowel while the inner one would defer the release of the protein with respect to the adjuvant. The interlayers formed from camostat mesilate or sodium glycocholate were also obtained by aqueous spray-coating with a 3.3% w/v solution of each compound containing 0.3% w/v of PEG 400 under the same process conditions employed for the HPMC coating. In place of tablets, hard- and soft-gelatin capsule cores were used as alternative substrates for the aqueous spray-coating procedure developed (Sangalli et al., 2009). An attentive set-up of the operating parameters was preliminarily required to prevent the sticking and shrinking of moisture- and heat-sensitive gelatin shells exposed to the nebulized water solution and warm inlet air flow. Particularly, an adequate balance between inlet air temperature and spray rate had to be established in the initial phases of the process in order to reach a product temperature range (40–50 ◦ C) that would enable a rapid solvent evaporation without altering the original volume of the units. As a result, the manufacturing of coated capsules having appropriate physico-technological characteristics and the pursued in vitro as well as in vivo release behaviour was feasible with no need for applying any protective sub-coating. The use of innovative coating techniques, such as primarily film-coating by tangential-spray rotary fluid bed and powder-layering, was also explored with the aim of improving the yield and reducing the duration of the manufacturing process as performed by top-spray fluid bed. Interestingly, the process time necessary to reach a Methocel® E50 coating level of approximately 50 mg/cm2 by fluid bed was more than halved in the case of tangential-spray film-coating in a rotary fluid bed, and shortened of approximately 75% with powderlayering, wherein the coating material is deposited in powder form and only limited amounts of liquids are employed for binding purposes. In order to enhance the efficiency of the HPMC coating in delaying drug release, thus reducing its thickness to values possibly complying with the size range of multiple units, a flexible and increasingly permeable outer film based on Eudragit® NE 30D and Explotab® V17, added as a pore-former, was applied (Maroni et al., 2013). Such a film, intended to initially hinder the penetration of water into the HPMC layer without affecting the relevant swelling, was obtained by aqueous spray-coating in a bottom-spray fluid bed. As an alternative to coating, the preparation of swellable/ erodible pulsatile release systems was recently accomplished by injection-moulding, which allowed functional shells to be manufactured from a hydrophilic cellulose ether with thermoplastic behaviour (HPC) (Gazzaniga et al., 2011; Zema et al., 2012, 2013). Such shells may serve as pre-formed “coatings” available for filling with a variety of drug formulations (solid, semisolid, liquid or multiparticulate). The choice of a device with appropriate capacity and thickness characteristics would depend on the apparent volume of the formulation to be conveyed and on the extent of delay that is sought. In a particular instance (Time Clock® system), the coating layer of an erodible pulsatile delivery system was based on hydrophobic materials of natural origin, such as carnauba wax and beeswax, mixed with the surfactant polyethylene sorbitan monooleate and a low-viscosity HPMC as the binder (Pozzi et al., 1994). Tablets of 100 mg containing either salbutamol sulphate or samarium oxide for ␥-scintigraphic imaging were coated in a fluid bed with a water dispersion of 5% w/w of waxes, 1% w/w of HPMC and 0.5% of surfactant. In an attempt to improve the nebulization and coalescence of the lipophilic solid particles though their melting and emulsification, the inlet air temperature was fixed at a relatively elevated value (75 ◦ C), which would restrict the applicability of this

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Fig. 6. Mean plasma concentration profiles of salbutamol from uncoated tablet cores ( ) and Time Clock® delivery systems ( ) with a 45% weight gain (6 fasting volunteers). Reproduced with permission from Pozzi et al. (1994).

technology to heat-stable bioactive compounds. Prior to release, coated systems with theoretical weight gains of 45 and 60% gave rise to reproducible in vitro and in vivo delays that were dependent on the coating level. Such delays were due to a progressive dispersion of the insoluble coating agents into the aqueous fluids. The performance observed by ␥-scintigraphy (6 volunteers) was unaffected by food intake. Moreover, the average salbutamol plasma concentration vs. time profile in the fasted state was consistent with that relevant to an immediate-release product in terms of AUC0–∞ and Cmax (Fig. 6). A correlation between in vitro and in vivo data could be found when increasing the viscosity of the test medium up to 120 cps by the addition of HPC, which resulted in extended in vitro lag phases. In an enteric-coated configuration, placebo and 5-aminosalicylic acid (5-ASA) Time Clock® systems with a 35% weight gain were demonstrated suitable for time-based colon delivery, as highlighted by ␥-scintigraphy and pharmacoscintigraphy (Wilding et al., 1994; Steed et al., 1997). 4. Permeable film coatings Permeable coatings are insoluble polymeric films through which the aqueous medium penetrates into a tablet core and the drug diffuses outwards after being dissolved. The duration of the lag phase depends on the thickness and composition of the functional film, which primarily affect the time required for water permeation. As opposed to rupturable coats, permeable membranes withstand disruption because of appropriate inherent permeability and mechanical resistance characteristics coupled with the use of inner formulations that would undergo no major expansion on water inflow. For this reason, the systems provided with such coatings generally perform as prolonged-release reservoir devices adapted for pulsatile delivery through the proper exploitation of the initial silent phase that is necessarily entailed by their particular design concept. With the aim of enhancing the rate of release after the lag phase, succinic acid and Eudragit® RS were combined within a peculiar multiple-unit formulation, the Sigmoidal Release System (SRS), after finding that organic acids added to the dissolution medium could expedite the drug liberation from pellets

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coated with such a polymer. SRS consisted of nonpareil seeds (24–32 mesh) loaded with the active ingredient in admixture with succinic acid and film-coated with the polymethacrylate by a fluid bed centrifugal granulator (coating formula: Eudragit® RS 30D/talc/TEC/water 39.0:5.7:1.2:54.1, operating temperature 50 ◦ C and oven-curing at 60 ◦ C for 16 h) (Narisawa et al., 1994, 1995, 1996). Following dissolution of the acid within the intermediate layer, its ionized and non-ionized forms would weakly interact with the positively-charged quaternary ammonium groups and the hydrophobic chains of Eudragit® RS, respectively, thus resulting in a looser polymer network with higher permeability. The overall time lapse required for the permeation of water across the outer coating, the consequent solubilization of succinic acid and, ultimately, the modification of the film permeability would account for the lag phase prior to release. A pulsatile release behaviour was obtained from SRS systems based on model drugs with differing solubility properties. The delays observed increased in duration as a function of the Eudragit® RS coating level (weight gains in the 10–80% range). Irrespective of the latter, they were followed by a relatively short diffusive phase and, in turn, by the fast release of most of the drug content, thus giving rise to the typical sigmoidal pattern of this delivery system. Pharmacokinetic studies confirmed the occurrence of a pulsatile release performance after administration of SRS to beagle dogs, pointing out a good agreement between in vitro and in vivo lag times. However, the slightly soluble model drug theophylline failed to be released quantitatively, which was attributed to dissolution constraints in the distal intestine. Because the use of dissolution media with increased osmolality turned out to extend the delay phase and slow down the drug liberation in vitro, osmotic pumping was hypothesized to be operating as an additional release mechanism (Narisawa et al., 1997). Another multiple-unit formulation based on Eudragit® RS was prepared by coating 1 kg of nonpareils (0.71–0.85 mm) with diltiazem hydrochloride (10% w/w in 95% w/v ethanol) and, successively, with Eudragit® RS 30D containing talc, TEC and polysorbate 80 as anti-tacking, plasticizing and solubilizing additives, respectively, up to weight gains of 5–12.5% (Kao et al., 1997). Although differing formulas were employed, the overall solid content of the polymethacrylate dispersions was maintained at 30% w/w. Both coating steps were carried out in a bottom-spray fluid bed under unchanged operating conditions. Finally, curing of two sub-batches was performed in a ventilated oven at 50 ◦ C for either 6 or 12 h in order to evaluate its influence on the release behaviour. This system proved able to delay the liberation of the model drug diltiazem hydrochloride as a function of the coating level, amount of plasticizer and curing time. High and pH-independent water solubility was necessary for the drug to be released in a relatively rapid mode at the end of lag time. The Eudragit® RS-coated formulations were therefore proposed for delivering drugs with such characteristics to various sites within the GI tract. An alternative coating material suitable for the formation of swellable diffusive films was obtained by UV cross-linking of Eudragit® L, which would suppress the pH-dependent solubility of the polymer (Hartman Kok et al., 2000). In this respect, pellets in the 0.6–0.7 mm sieve fraction (300 g batches), prepared by a highshear mixer, were coated in a fluid bed apparatus under the product temperature of 40 ◦ C with Eudragit® L 30D-55 containing 10% (on dry polymer) of TEC, 10% of pentaerythritol triacrylate as the crosslinking agent and 5% of 2,2-dimethoxy-2-phenyl-acetophenone as the photo-initiator. The coated units were dried in an oven at 60 ◦ C for 12 h and then exposed, in a fluidized state, to a UV light source at the wavelength of 365 nm. By conductivity measurements it was shown that the release from pellets coated up to weight gains of 33–154%, corresponding to 55–154 ␮m thicknesses, was delayed for few minutes depending on the coating level and on the time of exposure to UV cross-linking.

5. Semipermeable film coatings Semipermeable coatings are intended as insoluble films selectively permeable to water and impermeable to solutes. Such films constitute the key element of osmotic pumps, wherein they are responsible for establishing an osmotic gradient between the outer fluid and an osmo-active core formulation thus resulting in water influx into the latter. When the hydrostatic pressure that consequently builds up within the device offsets the osmotic pressure, an aqueous solution or dispersion of the active ingredient is slowly pumped out through one or more micrometric passageways present in the outer membrane. Cellulose acetate and other cellulose acylate derivatives are most commonly employed as film-forming agents for semipermeable coats (Verma et al., 2002). Due to the hydrophobic properties of such materials, organic spray-coating would represent the manufacturing technique of choice. Semipermeable films may be up to 200–300 ␮m thick because of the need for withstanding the hydrostatic pressure that is generated as water enters the system (Santus and Baker, 1995). The delivery orifices must possess a precise size in order to ensure the appropriate mode of release. If they are to large, they would indeed allow the drug to be delivered via diffusion (Verma et al., 2002). On the other hand, too small holes would affect the hydrostatic pressure developed and, ultimately, the overall release performance. Semipermeable membranes are usually punched after application, mainly by laser drilling that enables large-scale processing rates. However, the production of indented cores, wherein the micro-pores would be formed because of incomplete coating on the indentation area, or of tablets that undergo press-coating by means of modified tableting machines has also been described. Because of the time water takes for penetration into the core and dissolution of its components, the onset of release from osmotic pumps is necessarily delayed to an extent that depends on the nature and thickness of the semipermeable film. Although osmotic systems are mostly employed as prolonged release platforms, which is the use they were originally devised for, the occurrence of such a delay was exploited to design a verapamil hydrochloride formulation (once-a-day Controlled-Onset Extended-Release, COER-24) specially intended to address the chronotherapeutic needs of hypertension and angina pectoris (White et al., 1995; Gupta et al., 1996). Based on the OROS® Push-PullTM technology, this delivery system consisted in a two-layer osmotic tablet including a drug compartment and a push polymeric compartment (Prisant, 2001). The tablet core was surrounded by an inner hydrophilic cellulosic coating, aimed at further extending the lag phase, and an outer semipermeable film with orifices connecting the drug compartment with the exterior. On water inflow, the swellable polymer compartment expanded and pumped the drug out of the device. The release of the calcium channel blocker started after a delay period of 4–5 h and was prolonged over time, showing an agreement between in vitro and in vivo data (Gupta et al., 1996). Administered at bedtime to the patients, COER-24 proved able to match the circadian fluctuation patterns of blood pressure and heart rate, which typically exhibit peak values in the early morning hours (White et al., 1995). Osmotic formulations with a programmed delay phase were also proposed for time-based colonic release, the Osmet system being the first example (Chacko et al., 1990). Its ability to target the large bowel region was demonstrated in fasted and fed volunteers by ␥-scintigraphy. However, owing to the considerable size (25 mm × 7 mm) and limited capacity (200 ␮l), such a device was more suited for research rather than therapeutic use. With the aim of overcoming the inherent limitations of Osmet, a further osmotic colon delivery system was devised (OROS® -CT) based on the OROS® Push-PullTM technology (Verma et al., 2002). The

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OROS® -CT consisted of a hard-gelatin capsule containing 5 or 6 gastroresistant osmotic pumps of appropriate size. In contact with aqueous fluids the gelatin shell rapidly dissolved thus delivering its contents to the GI tract. The osmotic mechanism started operating after the sub-units left the stomach and the enteric film underwent pH-dependent dissolution. In a different time-dependent osmotic formulation for colon delivery, a metoprolol tartrate tablet was press-coated with a polysaccharide mixture and then spraycoated with a solution of cellulose acetate in methylene chloride and methanol containing HPMC and PEG as the pore-former and the plasticizer, respectively (Quadros et al., 1995). The hydrophilic polymer barrier interposed between the core and the semipermeable film was expected to prolong the lag phase duration. An orifice of 1 mm diameter was formed in the cellulose acetate membrane by means of a precision drill press. In agreement with the time-based approach to colon delivery, a gastroresistant coat was externally applied. The resulting system was able to delay the in vitro liberation of metoprolol for approximately 6 h (2 and 4 h in gastric and intestinal simulated fluid, respectively), which was confirmed by a pharmacokinetic study in fasted dogs. After the lag phase, zeroorder release was observed in vitro and in vivo. 6. Conclusions Oral pulsatile delivery systems have been attracting a growing research interest mainly because of their suitability for the chronotherapy of widespread diseases with typical night or early-morning symptoms. In most cases, they consist in coated formulations wherein an inner core containing the drug molecule is surrounded by a polymeric layer able to delay the onset of its release. Although a number of press-coated systems have been described, pulsatile delivery devices based on functional coats have primarily been attained by spray-coating. Thanks to the marked inherent versatility, such a technique has proved feasible in the production of the vast majority of these dosage forms despite the use of coating polymers with physico-chemical properties in a wide range. Moreover, it ensures flexibility in the selection of the thickness and composition of the films, relatively fast processing, scalability and ability to yield high-quality coatings. Therefore, the technical advances that may be expected in the field of oral pulsatile delivery systems, at least in the short term, will reasonably depend on progress made in spray-coating, provided that the utilization of organic solvents, largely employed until recent times, is properly restrained. The environmental and safety-related issues raised by the latter would indeed result in augmented costs and may fail to meet the increasingly tough requirements set for pharmaceutical production. The future evolution of coated dosage forms, however, may also rely on alternative techniques such as powder layering, three-dimensional printing and injection-moulding, along with those technologies limiting or avoiding the use of solvents that need to be removed through time- and energy-consuming drying steps. References Bianchini, R., Bruni, G., Gazzaniga, A., Vecchio, C., 1992. Influence of extrusionspheronization processing on the physical properties of d-Indobufen pellets containing pH adjusters. Drug Dev. Ind. Pharm. 18, 1485–1503. Bianchini, R., Bruni, G., Gazzaniga, A., Vecchio, C., 1993. d-Indobufen extendedrelease pellets prepared by coating with aqueous polymer dispersions. Drug Dev. Ind. Pharm. 19, 2021–2041. Bourgeois, S., Harvey, R., Fattal, E., 2005. Polymer colon drug delivery systems and their application to peptides, proteins, and nucleic acids. Am. J. Drug Deliv. 3, 171–204. Bussemer, T., Bodmeier, R., 2003. Formulation parameters affecting the performance of coated gelatin capsules with pulsatile release profiles. Int. J. Pharm. 267, 59–68. Bussemer, T., Dashevsky, A., Bodmeier, R., 2003a. A pulsatile drug delivery system based on rupturable coated hard gelatin capsules. J. Control. Release 93, 331–339.

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Bussemer, T., Otto, I., Bodmeier, R., 2001. Pulsatile drug-delivery systems. Crit. Rev. Ther. Drug Carrier Syst. 18, 433–458. Bussemer, T., Peppas, N.A., Bodmeier, R., 2003b. Evaluation of the swelling, hydration and rupturing properties of the swelling layer of a rupturable pulsatile drug delivery system. Eur. J. Pharm. Biopharm. 56, 261–270. Bussemer, T., Peppas, N.A., Bodmeier, R., 2003c. Time-dependent mechanical properties of polymeric coatings used in rupturable pulsatile release dosage forms. Drug Dev. Ind. Pharm. 29, 623–630. Chacko, A., Szaz, K.F., Howard, J., Cummings, J.H., 1990. Non-invasive method for delivery of tracer substances or small quantities of other materials to the colon. Gut 31, 106–110. Conte, U., Colombo, P., La Manna, A., Gazzaniga, A., Sangalli, M.E., Giunchedi, P., 1989. A new ibuprofen pulsed release oral dosage form. Drug Dev. Ind. Pharm. 15, 2583–2596. Conte, U., Giunchedi, P., Maggi, L., Sangalli, M.E., Gazzaniga, A., Colombo, P., La Manna, A., 1992. 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