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Controlling Drug Release Through Osmotic Systems
Controlled-release technologies play a vital role in today's current healthcare market, given the strong emphasis placed on product value and cost effectiveness. Treatment of many diseases requires a dosage regimen that delivers acceptable therapeutic concentrations of the drug at the site of action, which can be attained immediately and then constantly maintained over the desired duration of treatment (8). Controlled-release drug delivery offers solutions to conventional problems of drug administration by regulating the patterns of drug release and absorption as well as the localization of therapeutic agents. The maintenance of stable drug levels in the plasma over a defined and extended period, achieved with controlled-release systems, minimizes peak-to-trough variations of drug concentration in the systemic circulation and allows dosing frequency to be reduced, thereby improving patient compliance and overall clinical utility (1, 6). There are, however, potential disadvantages of controlled-release systems that should not be overlooked, such as dose dumping, possible toxicity or nonbiocompatibility of the materials used, undesirable by-products of degradation, higher manufacturing costs, and for implants, the requirement of surgical procedures to insert or remove these systems as well as possible patient discomfort from the delivery device (4, 5).
Osmotically controlled oral drug-delivery systems
Osmotically controlled oral drug-delivery systems have gained popularity because of the following advantages over other controlled-release
strategies (9, 11, 12):
Drugs with good water solubility may act as an osmotic agent or osmogen that draws water into the osmotic core. However, if the drug does not possess osmogenic properties, osmogenic salts (e.g., sodium chloride and potassium chloride) and sugars can be incorporated into the formulation (11). When selecting an osmogen, the two most important properties to consider are water solubility and osmotic activity.
The semipermeable membrane is an important component because it controls the rate of water influx into the drug core as well as retains water-soluble components within the core to create the osmotic pressure gradient that drives the osmotic system (9). The semipermeable membrane must possess certain performance criteria, such as sufficient wet strength and water permeability, and should be selectively permeable to water and biocompatible (13). Cellulose acetate, a water-insoluble film-forming polymer, is commonly used in osmotic systems. Release rate is affected by the molecular weight and acetyl content of the various grades of cellulose acetate. The semipermeable membrane usually contains a plasticizer, which moderates the permeability of the membrane, and in some cases, surfactants, flux regulators and pore-forming agents (11).
Formulation factors affecting drug release
Drug solubility: To achieve optimized drug release, the API for osmotic delivery should have sufficient water solubility, given that the release rate is directly proportional to the solubility of the API within the core. Drugs with extremes of solubility are generally poor candidates for osmotic delivery. There are, however, approaches to modify the solubility of such drugs within the core so that the desired release patterns can be attained.
For compounds with low solubility, solubilizing strategies can be employed; for example, by using alternative salt forms or cyclodextrins. Swellable polymers (e.g., vinyl acetate copolymer and polyethylene oxide) can also be added; the uniform swelling of these polymers facilitates drug release at a constant rate. Wicking agents help to increase the contact surface area of the drug with incoming fluids. The use of wicking agents can help enhance the rate of drug release from the orifice of the osmotic system (12, 14).
Osmotic pressure: The osmotic pressure gradient between the drug core of the osmotic system and the external environment is another important factor that controls drug release, with release rate being directly proportional to the osmotic pressure of the core. The simplest and most predictive way to achieve constant osmotic pressure would be to maintain a saturated solution of osmotic agent in the drug-core compartment (3).
Size of the delivery orifice: The size of the orifice must be within a certain range for controlled release. The typical range is 0.5 mm to 1.0 mm in diameter (9). For an optimal zero-order delivery profile, the cross-sectional area of the orifice must be small enough to minimize drug passage through the orifice but large enough to minimize the build-up of hydrostatic pressure within the osmotic system (14). The orifice can be created by using a mechanical drill or by laser drilling, which is now a well-established technology that offers reliability at low costs. Other methods to create an orifice are by indentation with modified punches that have a needle on the upper punch or by using leachable substances in the semipermeable coating (13).
Semipermeable membrane: Drug release rate is affected by the type and nature of the membrane-forming polymer used, membrane thickness, and the presence of other additives (e.g., type and nature of plasticizers used). Membrane permeability can be increased or decreased by the proper choice of membrane-forming polymers and other additives (3).
Osmotically controlled oral drug-delivery systems have come a long way. The first elementary osmotic pump, invented by Theeuwes, consists of a single compartment containing the drug and an osmotic agent surrounded by a semipermeable membrane (15, 16). Upon ingestion, water is drawn into the core through the semipermeable membrane to saturate the drug, which is then released in liquid form at a controlled rate through the orifice(s) (see Figure 1a).
The limitation of the elementary osmotic pump, however, is that it can only deliver water-soluble drugs. The design was further improved by Cortese and Theeuwes, resulting in the development of the push-pull osmotic pump, which is a bilayer tablet capable of delivering both highly and poorly soluble drugs (17). The upper layer (i.e., the drug layer, also known as the pull layer) consists of the drug and an osmotic agent while the lower layer (i.e., the push layer) consists of water-swellable polymers and osmotic agents (see Figure 1b). Both layers are coated with a semipermeable membrane that regulates water influx into the system. As water enters the tablet, pressure increases and the polymer swells to push against the drug layer, thereby releasing the drug solution or suspension through the laser-drilled orifice(s).
Zentner et al. reported on the development of the controlled-porosity osmotic pump in the mid-1980s. The controlled-porosity osmotic pump does not require a delivery orifice for drug release, hence eliminating the need for complicated laser-drilling procedures (18, 19). It consists of the drug and an osmotic agent in a tablet core surrounded by a semipermeable coating membrane containing leachable pore-forming agents, which dissolves upon contact with water, forming pores through which the drug solution is pumped out. The rate of drug release is dependent on the thickness of the coating membrane, levels of leachable pore-forming agents, the amount of soluble components incorporated in the coating, drug solubility within the tablet core and the osmotic pressure differences across the membrane, but unaffected by pH and gastrointestinal motility (20–22).
In the mid-1990s, Herbig et al. described a new type of membrane coating for osmotic drug delivery. The new coating has an asymmetric structure, similar to asymmetric membranes made for reverse osmosis or ultrafiltration, in that the coating consists of a porous substrate with a thin outer skin (23). Asymmetric membranes are made from water-insoluble polymers (usually cellulose derivatives, such as cellulose acetate, ethyl cellulose and cellulose acetate butyrate) and pore-forming agents (e.g., glycerol, sorbitol, polyethylene glycol, polyglycolic acid and polylactic acid) using a phase-inversion process.
The use of asymmetric membrane coatings in osmotically controlled oral drug-delivery systems has increased in the past decade because of the advantages it offers. These benefits include: a higher rate of water influx, which facilitates osmotic delivery of poorly soluble drugs and enables higher release rates of such drugs; more controlled-release of freely soluble drugs; pH-independent release and minimized exposure to the gastrointestinal tract, which results in reduced gastric irritation and degradation of drugs (23–25).
Water permeability of the coating can be adjusted by controlling the membrane structure, thereby allowing control of release kinetics without altering the coating materials used or significantly varying the coating thickness. The porosity of the membrane can also be controlled to minimize the lag time that occurs before drug delivery begins (23).
Asymmetric membrane coatings can be applied on pharmaceutical tablets and capsules (23, 26, 27). The basic design of an asymmetric membrane capsule is similar to a hard-gelatin capsule, except that the shell contains pore-forming water-soluble additives, which dissolve after coming in contact with water, resulting in an in situ formation of a microporous structure (28, 29). A delivery orifice is, therefore, not required due to the in situ pore formation of the asymmetric membrane.
Asymmetric membrane capsules in application
Philip and Pathak developed a nondisintegrating, controlled-release, asymmetric membrane capsule of ketoprofen and evaluated the in vitro and in vivo correlation of the formulation (30). Ketoprofen is a nonsteroidal anti-inflammatory drug, used in treatment of rheumatoid arthritis, osteoarthritis, and musculoskeletal disorders. Multiple dosing is required to achieve and maintain therapeutic concentration because of its short half-life (4.2 h) and poor solubility. The asymmetric membrane capsule was made of ethyl cellulose and glycerol. Sodium chloride was used as an osmotic agent and citric acid as a solubilizer. The formulation provided controlled release of ketoprofen and the half-life of the drug was prolonged for more than 16 hours. In vivo pharmacokinetic studies showed excellent level A correlation (r2 > 0.99), demonstrating that the in vitro drug-release profile of ketoprofen from the asymmetric membrane capsule could be used to accurately predict the in vivo performance.
A double-membrane system of cefadroxil using ethyl cellulose as the inner membrane and cellulose acetate phthalate as the outer membrane has been developed by Philip et al. (31). The asymmetric membrane in a membrane capsule was prepared on glass pins using a two-step phase-inversion process. The first step was to form a nondisintegrating, asymmetric membrane capsule and the second step involved formation of a pH-sensitive, disintegrating, asymmetric membrane formed over the nondisintegrating membrane. This double-membrane formulation was able to delay the release of cefadroxil for the first two hours in the gastric medium and provide controlled release in the intestinal medium for an extended 12-hour period. Drug release was independent of pH and agitation intensity and followed zero-order release kinetics.
More recently, Garg et al. reported on the development of asymmetric membrane capsules with two compartments for simultaneous delivery of two poorly soluble antihypertensive drugs, atenolol and amlodipine (32). Scanning electron microscopy showed a dense outer region and a porous inner region of the asymmetric membrane before dissolution. Pore size of the outer and inner layers increased after dissolution. Buffering agents were used to increase the solubility of atenolol and amlodipine. The formulation followed zero-order release kinetics, which was not affected by the agitation intensity of the dissolution fluid. Drug release was dependent on the diffusion rate of the drug across the membrane and the osmotic pressure. The effect of membrane thickness on dissolution fluid entering the asymmetric membrane capsule showed that as the membrane thickness increased, the volume of dissolution fluid entering into the system decreased.
A gastroretentive asymmetric membrane capsule of famotidine was recently developed by Guan et al. (33). The drug is used for the treatment of duodenal, gastric and peptic ulcers. It has a relatively short half-life (3 h) and a low bioavailability (45–50%). Increasing gastric residence time would allow the drug to penetrate through the gastric mucus layer and produce a more pronounced effect. Polyethylene oxide was used as floating agent in the formulation. Pharmacokinetic studies in beagle dogs showed that the gastroretentive asymmetric membrane capsule displayed complete drug delivery with zero-order release kinetics and a 12-hour floating time. The system had the ability to prolong drug action, minimize dosing frequency, and reduce the average peak plasma concentration.
Chauhan et al. also reported on the design of a floating asymmetric membrane capsule for site-specific delivery of ranitidine in a controlled manner (34). Ranitidine was given 150 mg twice daily or 300 mg once daily as an oral dosage form. Dosing has to be increased to 150 mg 4–5 times a day for the treatment of endoscopically diagnosed erosive esophagitis. The conventional formulations could only inhibit acid secretion for up to 5 hours, hence, requiring frequent dosing, which would cause fluctuations of drug levels in the plasma. The aim was to develop a buoyant asymmetric capsule with density less than the gastric fluid for controlled release of ranitidine in the gastric cavity. The capsule shell was prepared by the phase-inversion process wherein the polymeric membrane was precipitated on glass pins by dipping them in a solution of cellulose acetate followed by quenching. The solubility of ranitidine was suppressed by the ion effect, using optimized coated sodium chloride crystals as a formulation component. Drug release with zero-order kinetics was achieved and the asymmetric membrane capsule demonstrated floating ability for up to 12 hours.
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