PIPAAm grafted gels showed rapid deswelling kinetics without formation of skin layer on surface. This is probably due to rapid dehydration of graft chains formed by hydrophobic aggregation on the three-dimensional cross-linked chains. The molecular weight of the graft chains had a significant effect on the gel deswelling kinetics and also on the drug release profiles, especially for large molecular weight drugs. ^[16-18] A similar rapid deswelling phase was achieved by incorporating poly(ethylene glycol) graft chains in PIPAAm cross linked hydrogels. ^[19] The low molecular weight compounds released immediately from conventional PIPPAm gels after a temperature increase, after which the release was terminated due to the formation of a dense impermeable skin layer on the surface. In comparison, 65% of the drug was released in one burst from free PIPPAm grafted hydrogels with a graft molecular weight of 9000, following the temperature increase. Graft-type gels with a molecular weight of 4000 showed oscillating drug release profiles. The release of high molecular weight compound from PIPPAm graft type gels was shown to burst after a temperature increase of 40 ºC, while release was suppressed from IGG 4000. The difference in drug release profiles for two graft-type gels is probably due to the different strengths of aggregation forces between the formed hydrophobic cores within the graft-type gels. The large molecular weight graft chains formed more hydrophobic cores within the gels upon the temperature increase, which induced rapid gel deswelling. In contrast, aggregation forces between graft chains in IGG 4000 were relatively weak, thus leading to the formation of skin layer on the gel surface, which limited the drug diffusion from gel interior. A similarly rapid deswelling phase was achieved by incorporating poly(ethylene glycol) (PEG) graft chains into PIPPAm cross-linked hydrogels. The introduction of PEG chains did not alter the transition temperature. ^[20] Yuk et al. designed temperature-sensitive drug delivery systems using a mixture of poly(ethylene oxide)-poly(propylene oxide)-poly(ethylene oxide) triblock copolymer (F-68) and poly vinyl alcohol (PVA). Change in the ratio of F-68/PVA could be used to manipulate the swelling transition of polymer complex gel. Pulsatile release of acetaminophen in response to change in temperature between 35 and 40ºC has observed. ^[21] US Patents 6733788 and 20020015712 describe a medical device containing thermo-sensitive cellulose gel structure, which can deliver the bioactive solute compounds to a target location in the body. The gel structure deswells at certain temperature and expels the biologically active solute with an increase in gel temperature. The loaded solute released in a relatively short period of time under the influence of increased temperature of the body.^[22,23] Thermo-responsive polymeric micelle systems constitute polymeric micelles whose properties and biological interests make them a most noteworthy candidate as drug carrier for the treatment of cancer.^[24] The polymeric micelle is composed of amphiphilic block copolymers exhibiting a hydrophobic core with a hydrophilic corona. Due to these unique characteristics, polymer micelles exhibit stealth characteristics and are not detected by the body defense system (reticuloendothelial system). Thus passive targeting could be achieved through enhanced permeation retention (EPR) effect of tumor sites.^[25] Okano and coworkers used an end functionalized PIPAAm to prepare block copolymers. Hydrophobic polymers, such as poly(butyl methacrylate) (PBMA), polystyrene (PSt), poly(lactic acid) (PLA) were used. ^[26-29] Block copolymers formed micellar structure (with core-shell structure) in aqueous solution below PIPAAm's transition temperature. The shell was made from thermo-responsive PIPAAm, while the core consisted of hydrophobic polymer aggregates of poly(butyl methacrylate) (PBMA). The hydrated PIPAAm corona acted as an inert material towards all biological entities below PIPAAm's LCST. However, upon temperature increase above 32° C hydrated PIPAAm chains became hydrophobic, due to dehydration of polymer chains, thus resulting in aggregation and precipitation. The hydrophobic anticancer drug, andriamycin, was incorporated in to PBMA micelle cores. At temperatures below PIPAAm's low crystalline solution temperature (LCST), drug release was at a minimum, with a value less than 10%. However upon temperature increase above PIPAAm's LCST, accelerated release of andriamycin was observed.

2. Chemical Stimuli Induced Pulsatile Systems:

(a) Glucose-Responsive Insulin Release Devices:

A number of stimuli responsive hydrogels have been developed for controlled release of insulin which respond to glucose concentration changes. One such system is based on glucose oxidase (GOD) catalysed glucose oxidation. Ishihara et al. reported two types of gel membrane systems GOD and nicotinamide-immobilized gel membranes. When glucose was oxidized by the immobilized GOD, the resulting hydrogen peroxide oxidized the nicotinamide units inducing positive charges. The gel membranes became hydrophilic and permeability get increased. ^[30] In other study, GOD-immobilized hydrogels were prepared from 2-hydroxy-ethyl methacrylate and N,N-dimethylaminoethyl methacrylate. Glucose in the medium was transformed to gluconic acid through a reaction mediated by hydrogel-immobilized GOD. Protonation of hydrogel amino groups lead to swelling of hydrogel membrane thereby induced this type of membrane, the insulin permeability through the hydrogels were found to be higher. ^{[31, 32]} Kim and co-workers have developed self-regulating insulin delivery systems using microcapsules containing the glucose-binding lectin, concanavalin A (Con A) and Con A-bound glucosylated insulin. Glucosylated insulin bound to Con A was released through exchange with external glucose, due to the difference in their binding constants direct injection of microcapsules into the peritoneal cavity may cause undesirable side-effects arising from the immune response toward Con A. Moreover, refilling of glucosylated insulin after exhausting initially bound insulin molecules is required. ^{[33, 34]} Another insulin drug delivery system is based on copolymer of acrylamide and allyl glucose. The side chain glucose units in the copolymer were bound to Con A. These hydrogels showed a glucose-responsive, sol–gel phase transition dependent upon the external glucose concentration. The nonlinear dependence of this sol–gel phase transition with regard to the glucose concentration was not only due to the increased binding affinity of allyl glucose to Con A compared to native glucose, but the cooperative interaction between glucose containing copolymer and Con A as well. ^[35,
36] Okano et al. focused on the boronic acid moiety, which is known to form reversible bonds with polyol compounds, including glucose. Water-soluble copolymers, containing phenylboronic acid side chains, have formed reversible complex gels with polyol compounds such as poly(vinyl alcohol) (PVA). ^[37,
38] Such complexes dissociated after the addition of glucose in a concentration dependent manner. Utilizing this polyol-binding characteristic of boronic acid, microgel beads (containing phenyl-boronic acid) with a diameter of 100–400 mm were prepared. Gluconic acid-modified insulin was bound to the boronic acid moiety of the beads. The bound insulin was released after substitution with external glucose molecules. The only disadvantage is that the system works at pH 8.5 and, therefore, much higher than physiological conditions. ^[39] Shiino et al. incorporated amino groups in the vicinity of the boronic acid in the gels. The lone electron pair of the amino nitrogen atom was coordinated with the boron atom in such a position as to allow the formation of stable complexes with the polyol at a pH lower than 8.5. Thus, glucose-sensitive insulin release was achieved at pH 7.4 using amine-containing phenyl-boronic acid copolymer gel beads. ^[40] Kataoka et al. recently developed glucose and thermo-responsive hydrogels using acrylamidophenylboronic acid and N-iso-propylacrylamide (IPAAm). The obtained gels, containing 10 mol% phenylboronic acid moieties, showed a transition temperature of 22 ºC in the absence of glucose. Below this temperature, the gels existed in a swollen state. The introduction of glucose to the medium altered the transition temperature of the gels in such a way that the transition temperature increased with increasing glucose concentration to reach 36 ºC at 5.0 g/ l glucose concentration. Boronic acid was in equilibrium between the undissociated (uncharged) and the dissociated (anionically charged) form. With increasing glucose concentration, the equilibrium shifted to increase the amount of dissociated boronate groups and the gels became more hydrophilic. Thus, the data indicated that the gel’s swelling / deswelling was regulated by glucose concentration at a fixed temperature. ^[41]

(b) Inflammation-Induced Pulsatile Release:

When human beings receive physical or chemical stress, such as injury, broken bones, etc., inflammation reactions take place at the injured sites. At the inflammatory sites, inflammation responsive phagocytic cells, such as macrophages and polymorphonuclear cells, play a role in the healing process of the injury. During inflammation, hydroxyl radicals (^.OH) are produced from these inflammation responsive cells. Yui and co-workers focused on the inflammatory induced hydroxyl radicals and designed drug delivery systems, which responded to the hydroxyl radicals and degraded in a limited manner. They used hyaluronic acid (HA), a linear mucopolysaccharide composed of repeating disaccharide subunits of N-acetyl-D-glucosamine and D-guluronic acid. In the body, HA is mainly degraded either by a specific enzyme, hyaluronidase, or hydroxyl radicals. Degradation of HA via the hyaluronidase is very low in a normal state of health. Degradation via hydroxyl radicals however, is usually dominant and rapid when HA is injected at inflammatory sites. Thus they prepared cross-linked HA with ethylene glycol diglycidylether or polyglycerol polyglycidylether. These HA gels degraded only when the hydroxyl radicals were generated through the Fenton reaction between Fe²⁺ ions and hydrogen peroxide in vitro. Thus, a surface erosion type of degradation was achieved. When microspheres were incorporated in HA hydrogels as a model drug, these microspheres were released only when hydroxyl radicals induced HA gel degradation. The microsphere release was regulated by the surface erosion type of degradation. Furthermore, in vivo tests of HA hydrogel degradation showed that HA gels were degraded only when inflammation at the implanted site was induced by surgical incision. Control HA gels implanted in the animals were found to be relatively stable over a period of 100 days. Thus, it is possible to treat patients with inflammatory diseases, such as rheumatoid arthritis, using anti-inflammatory drug incorporated HA gels as new implantable drug delivery systems.^{[42, 43]}

(c) Drug Release from Intelligent Gels Responding To Antibody Concentration:

There are numerous kinds of bioactive compounds which exist in the body. Recently, novel gels were developed which responded to the change in concentration of bioactive compounds to alter their swelling / deswelling characteristics. Miyata and co-workers focused on the introduction of stimuli-responsive cross-linking structures into hydrogels. Special attention was given to antigen antibody complex formation as the cross linking units in the gel, because specific antigen recognition of an antibody can provide the basis for a new device fabrication. Utilizing the difference in association constants between polymerized antibodies and naturally derived antibodies towards specific antigens, reversible gel swelling /deswelling and drug permeation changes occurred. Thus, biological stimuli-responsive hydrogels were created. ^{[44, 45]}

3. pH Sensitive Pulsatile Release:

It is a widely accepted and versatile approach to design chronotropic systems to attain specified lag time prior to drug release by using pH dependent polymers. These can be single unit or multiparticulate dosage forms with reliable and predictable drug release profile. These type of systems possess the advantage of fact that there exists different pH environment at different parts of gastrointestinal tract. Hence by employing pH dependent polymers targeting at specific site of gastrointestinal tract is possible as well as a desired lag time can be achieved due to dependency of polymer solubility only at a particular pH of gastrointestinal tract. Examples of pH dependent polymers include copolymers of methacrylic acid (various grades of Eudragit), phthalates, carboxy methyl cellulose etc. These polymers are utilized for enteric coating to protect the drug from degradation in upper G.I.T and attain drug release at specific part of intestine (according to solubility of polymer at particular pH and specific site of intestine) after a predetermined lag time. A number of chronotropic systems have been developed and marketed for chronotherapy utilizing pH dependent polymers for asthma, angina, rheumatoid arthritis, cancer, diabetes, ulcer etc. ^[46]

(a) pH Dependent Colonic DDS:

A progressive pH increase is observed from the stomach to the rectum. pH of stomach is 2-3 (higher after eating), 6.5-7 in the small intestine and 7-8 in the large intestine. pH sensitive polymers especially those that contain carboxyl groups which make them insoluble at low pH values and soluble at higher pH values are used for colon targeting of drug. Eudragit L (soluble at pH>6 specifically used as enteric polymer) and Eudragit S and FS (dissolve at pH> 7, specifically dissolve in colon/rectum) and their combinations are used for colon targeting. Tablets, minitablets or pellets containing drug are coated with these polymer for achieving colon delivery. pH-based systems are commercially available for number of drugs.

(b) pH and Time Dependent Colonic Systems:

Combination of enteric polymer with delayed release component that relies on the passage of time, lead to feasibility of target the drug in the terminal ileum and colon. These systems are comprised of (i) An outermost coating of an enteric polymer such as Eudragit L, S or FS that ensures dosage transit from stomach to some distance into small intestine. (ii) A second barrier coating of a pH-independent polymer such Eudragit RS or ethyl cellulose that delays drug release for several hours. (iii) An innermost drug layered nonpareil seed. One example of this system is colon targeted delivery capsule. The system contains an organic acid that is filled in a hard gelatin capsule as a pH- adjusting agent together with the drug substance. This capsule is then coated with a three layered consisting of an acid –soluble layer, a hydrophilic layer and an outermost enteric layer. After ingestion of the capsule, these layers prevent drug release until the environment pH inside the capsule decreases by dissolution of the organic acid, upon which the enclosed drug is quickly released. Therefore, the onset time of drug release is controlled by the thickness of the acid-soluble layer. In fact, capsule disintegration does not start until 5h after gastric emptying regardless of whether the formulation is administered to fasted or fed subjects.

(c) pH and Bacterial Enzyme–Dependent Colonic DDS:

A unique technology that combines the advantages of bacterial enzyme-degradable polysaccharides and pH- sensitive polymers is the CODES system. The outer coating is composed of a standard enteric polymer such as Eudragit L which dissolves after the transit of dosage form into the small intestine and exposes the undercoating, which is composed of Eudragit E (acid soluble). This coating will not dissolve in the environment of the small or large intestine. The undercoating permits lactulose to be released into the environment adjacent to the tablet. This disaccharide is metabolised to short chain fatty acids, which lower the local pH to point where the Eudragit E dissolves. This final dissolution step exposes the core of the tablet permitting drug dissolution to occur. ^[47]

B. EXTERNAL STIMULI INDUCED PULSATILE SYSTEM:

1. Electric Stimuli-Responsive Pulsatile Release:

An electric field as an external stimulus provide several advantages such as the availability of equipment, which allows precise control with regards to the magnitude of current, duration of electric pulses, interval between pulses etc. Electrically responsive delivery systems are prepared from polyelectrolytes (polymers which contain relatively high concentration of ionisable groups along the backbone chain) and are thus, pH-responsive as well as electro-responsive. Under the influence of electric field, electro-responsive hydrogels generally deswell or bend, depending on the shape of the gel lies parallel to the electrodes whereas deswelling occurs when the hydrogel lies perpendicular to the electrodes. Synthetic as well as naturally occurring polymers, separately or in combination, have been used. Examples of naturally occurring polymers include hyaluronic acid, chondroitin sulphate, agarose, carbomer, xanthan gum and calcium alginate. The synthetic polymers are generally acrylate and methacrylate derivatives such as partially hydrolyzed polyacrylamide, polydimethylaminopropyl acrylamide. Many of these gels are prepared by either cross-linking the water-soluble polymers using radiation or chemical agents such as Ca²⁺ions, ethylene diglycidylether, N,N´-methylene bisacrylamide, ethylene glycol dimethacrylate or by free-radical polymerization of monomers. Complex multi-component gels or interpenetrating networks have been prepared in order to enhance the gel’s electro responsiveness. ^[48] Several approaches have been reported describing the preparation of electric stimuli-responsive drug delivery systems using hydrogels. Kishi et al. developed an electric stimulus induced drug release system based on swelling / deswelling Characteristics of polyelectrolyte hydrogels. These gels exhibited reversible swelling / shrinking behaviour in response to on–off switching of an electric stimulus. Poly(sodium acrylate) microparticulate gels containing pilocarpine as a model drug were prepared. Pilocarpine release from microparticulate gel beads was found to be increased with applied electric stimuli. However, because the matrix itself showed higher swelling in the medium, pilocarpine release was not terminated after removal of the electric current. Thus, on–off release regulation of drugs could not be realized with this system. ^[49] Kwon and co-workers used cross-linked poly(2-acrylamide-2-methylpropanesulfonic acid-co-butyl methacrylate) (P(AMPS-co-BMA)) hydrogels for electric stimuli-induced drug delivery systems. Positively charged edrophonium chloride was incorporated within negatively charged (P (AMPS-co-BMA)) hydrogels via an electrostatic interaction and investigated the drug release from the hydrogels in an electric field. By applying electric field, ion-exchange between edrophonium ions and protons commenced at the cathode, resulting in a rapid drug release from the hydrogels. The drug release rate was increased by increasing applied voltage in a dose-dependent manner. Interestingly, complete on–off drug release regulation was achieved with this system, as no drug release was apparent without the application of an electric field. Electric stimuli induced release of, the neutral drug, hydrocortisone from (P (AMPS-co-BMA)) hydrogels via a diffusion mechanism was developed and the release of hydrocortisone was found to show a small pulse when an electric current was applied. These pulses, however, may be attributed to the bulk squeezing caused by anisotropic deswelling under the electric field. This hydrogel system may be a potential candidate for the delivery of positively charged drugs in conjunction with iontophoresis. ^[50–52] Kwon et al. synthesized a different type of electroresponsive hydrogel. The gels were formed when two aqueous polymer solutions were mixed and a complex was formed due to hydrogen bonding or ionic bonding between the polymers over a certain pH range. When the polycationic polyamine solution was mixed with the polyanionic heparin solution, a complex was formed via ionic bonding between the positively charged NH³⁺ groups in polyallylamine and the COO^- and SO^3- groups in heparin, over the broad pH range 3-10. They also achieved ‘on-off’ drug release utilizing a polyethyloxazoline (PEO) and polymethacrylic acid (PMAA) from solid complexes by hydrogen bonding below pH 5, but the complexes dissolve above pH 5.4. When an electric current was applied through the disk-shaped matrix of the complex in a saline solution, the matrix dissolved at the surface facing the cathode because local pH increase near the cathode and resultant hydrogen bonding was disrupted. The release pattern of heparin from the polyelectrolyte complex exhibited complete on–off profile in response to on–off stimuli of electric current. Insulin released from the matrix with dissolution in response to application of the electric current. ^[53] Feedback-controlled drug delivery of an ionic drug was investigated based on the electrophoresis. In this system, drug release was regulated in either a positive or a negative manner by appropriate control of the electric field between a pair of platinum electrodes placed on either side of rate controlling polymeric hydrogel membrane. Control of the electric field was carried out by the microprocessor, which received signals from the biosensor element. ^[54] Kiser et al. designed lipid-coated microgels for the triggered release of drugs. Ionic microgels were synthesized from the monomers of methylenebisacrylamide (MBAM), methylacrylic acid (MAA) and 4-nitrophenyl methacrylate (NPMA) and coated with a lipid bilayer. The release of drug was triggered from the gels using either lipid solubilising surfactants or electroporation. The swelling and release of drugs in three stages: i) the permeability of the membrane might be sufficiently compromised (e.g. by electroporation or membrane dissolution or other permeabilizing species), but only to an extent that allows proton efflux from the microgel and a sodium ion influx into the gel particle; ii) microgel begins to swell due to occurrence of exchange process, allowing additional ions to be transported across the membrane and so that disruption of membranes causes uncoating of microgel; and iii) drug is exchanged from the hydrogel by Na⁺ ions and diffuse down its concentration gradient out of the expanded polymer network into the surrounding medium over a period of time, resulting in a triggered release.^[55] Electronic microelectromechanical devices was manufactured using standard microfabrication techniques that are used to create silicon chips for computers and they often have moving parts or components that enable some physical or analytical function to be performed by the device. such devices are commonly referred to as “microelectromechanical systems “(MEMS). MEMS are found in ink-jet printers, automotive applications, and microtube engines used in the aerospace industry. MEMS for biological applications are classified as either microfluidic devices or nonmicrofluidic and release multiple chemical substances on demand by a mechanism devoid of moving its parts. ^{[56, 57]} The usage of MEMS is triggered, particularly pulsatile delivery of drugs represents a new area of study that is to be explored. One straightforward approach to achieve this drug reservoir is the fabrication of silicon microparticles that contain an internal reservoir loaded with drug. These devices could be used for oral drug delivery, with release of the drug triggered by binding of a surface-functionalized molecule to cells in the digestive tract.^[58] The completely implantable minipump made by Minimed^R, has also a pulsatile, radio controlled injection rate through a catheter into the intraperitoneal region. It was found that patients with the implantable pump did not differ from control subjects on any measure of psychosocial function but that pump users monitored their blood glucose levels more frequently and had lower average blood glucose levels.^[59] Even though this type of device may improve patient’s mobility and reduce infections by eliminating transcutaneous catheters, they may still be hampered by their size, cost, ability to deliver only drugs in solution and the limited stability of some drugs in solution at 37ºC. Ikemoto et al. developed a stepmotor micropump for the injection of nanolitre(nl) volumes of drug D-amphetamine solution into discrete brain regions of freely moving rats was and it found to be well tolerated. This micropump delivered a reliable volume of 50 nl per infusion over an hour at a rate of one infusion per minute. ^[60] Microchip delivery devices, described in US Patents 579898, US2006123861 and 0121486 provide a means to control both the rate and time of release of variety of bioactive molecules, in either a continuous or pulsatile manner. Hundreds or thousands of reservoirs can be fabricated on a single microchip. The molecules to be delivered were inserted into the reservoirs by injection or spin coating methods in their pure form or in a release system. The release systems include polymers and polymeric matrices, non-polymeric matrices, and other common excipient or diluents. The physical properties of the release system control the rate of the release of molecules. Release from an active device is controlled by a pre-programmed microprocessor, remote control, or by biosensors. ^[61-63] US Patents 20060057737A1 and 20060178655A1 described a device and method for the time and/ or rate controlled release of one or more drugs. The device includes implantable microchips which have reservoirs containing the bioactive compounds for the triggered or controlled release. The microchip devices include i) a substrate, ii) at least two reservoirs in the substrate containing the molecules for release, and iii) a reservoir cap positioned on, or within a portion of the reservoir and over the molecules, so that the molecules are released from the device by diffusion through or upon disintegration or rupture of the reservoir caps in controlled manner. Each of the reservoirs of a single microchip can contain different molecules and/or different amounts and concentrations, which can be released independently. The filled reservoirs can be capped with materials that passively or actively disintegrate. Passive release reservoir caps can be fabricated using materials that allow the molecules to diffuse passively out of the reservoir over certain time. Active release reservoir caps can be fabricated using materials that disintegrate upon application of electrical, mechanical, or thermal energy. Release from an active device is also controlled by a preprogrammed microprocessor, remote control, or by biosensors. ^[64,
65] US Patents 20060121486A1 and 20060100608A1 described the device includes a reservoir cap formed of an electrically conductive material, which prevents the reservoir contents from passing out from the device and prevents exposure of the reservoir contents to molecules outside of the device; electrical input and output leads connected to the reservoir cap, such that upon application of an electrical current through the reservoir cap, via the input and output lead, the reservoir cap ruptures to release or expose the reservoir contents.^[66,67] US Patent 20060105275A1 describes modified fabrication methods and structures for micro-reservoir devices. The multi-reservoir device comprising i) patterning one or more photoresist layers on a substrate; ii) depositing onto the substrate at least one metal layer by a sputtering process to form a plurality of reservoir caps and conductive traces; iii) removing the photoresist layers using a lift off process; iv) forming a plurality of reservoirs in the susbtrate; v) loading each reservoir with bioactive molecules; and vi) sealing of each reservoir. The reservoir cap comprises a first conductive material coated with one or more protective metal films, such as gold, platinum, silicon carbide, silicon dioxide and platinum silicide.^[68]

2. Magnetically Induced Pulsatile Release:

The use of an oscillating magnetic field to modulate the rates of drug release from polymer matrix was one of the old methodologies. Magnetic carriers receive their magnetic response to a magnetic field from incorporated materials such as magnetite, iron, nickel, cobalt etc.^[69] For biomedical applications, magnetic carriers must be water-based, biocompatible, non-toxic and non-immunogenic. Magnetic steel beads were embedded in ethylene vinyl acetate copolymer (EVAc) matrix that was loaded with bovine serum albumin (model drug) and increased rates of drug release were obtained in the presence of an oscillating magnetic field. ^[70] During exposure to the magnetic field, the beads oscillate within the matrix, alternatively creating compressive and tensile forces. This in turn acts as a pump to push an increased amount of the drug molecule out of the matrix. Co-polymers with a higher Young's modulus were more resistant to the induced motion of steel beads, and consequently the magnetic field has less effect on the rate of drug release from these materials.^[71] Saslawski et al. developed different formulations for in vitro magnetically triggered delivery of insulin based on alginate microspheres. In an experiment, ferrite microparticles (1 mm) and insulin powder were dispersed in sodium alginate aqueous solution. The ferrite-insulin alginate suspension was later dropped in aqueous calcium chloride solution which causes the formation of cross linked alginate spheres, which were further cross linked with aqueous solution of poly(L-lysine) or poly(ethylene imine). The magnetic field characteristics due to the ferrite microparticles and the mechanical properties of the polymer matrices were supposed to play role in controlling the release rates of insulin from the system.^[72] Babincova et al. had developed magnetoliposomes for triggered release of drug. In their delivery systems, they entrapped dextran-megnetite and model drug, 6-carboxyfluorscein in the liposomes and used laser as triggering the release of model agent. The magnetite absorbs the laser light energy to heat the lipid bilayer above the gel-liquid crystal phase transition temperature (Tc), which is 41 ºC for dipalmitoylphosphatidylcholine. Liposomes made from this lipid release their content as soon as temperature is reached to this level. They have also suggested that the absorption of laser energy by magnetite particles provide a means for much localized heating and controlled release of liposome with a single laser pulse. This may have potential applications for selective drug delivery especially to the eye and skin e.g. a number of products patented based on magnetically induced drug delivery system. ^[73] US Patent 20016251365 describes about the magentosomes comprising magnetic monocrystals having a maximum diameter of 45 nm surrounded by a phospholipids membrane developed from triggered release in the body. The membrane consists of phophotidylethanolamine, phosphotidyl glycerol and phosphotidyl choline containing mainly the fatty acids such as palmitic acid, palmitoloeic and oleic acid. These magnetosomes having cationic charge increase the probability that the antibodies and therapeutic agents can be correctly bound to them as a potential carrier for antibodies. ^[74] US Patent 20036514481B1 describes about the nanosized particles termed as “nanoclinics” or “nanoparticles” or “nanobubbles”, prepared from iron oxide particles using a reverse micelle colloidal reaction. A tracking agent such as a two photon is attached to the surface of the particles to track the nanoparticles using two-photon laser scanning microscopy. Silica (sodium salt) is added to perform the silica shell. A carbon spacer is then attached to the silica surface to reduce the steric hindrance for the binding of the targeting agent to spacer. These nanoparticles can be used for lysis of cells, such as for the treatment of cancer. For the lysis of tumor cells, the targeting agent is selected based on the molecules on the surface of the tumor cells. Following binding and/ or internalization of the nanoparticles by the tumor cells, the particles are exposed to DC magnetic field. Dc magnetic field can be obtained by standard Magnetic Resonance Imaging (MRI) equipment which typically has a magnetic filed in the range of 0.1 to 5 Tesla. ^[75] US Patent 2006997863B2 provides a treatment method for hyperthermia that involves the administration of a magnetic material composition, which contains single-domain magnetic particles attached to a target-specific ligand, to a patient and the application of an alternating magnetic field to inductively heat the magnetic material composition, which cause the triggered release of therapeutic agents at the target tumor or cancerous cells.^[76]

APPLICATIONS:

(a) Parenteral Delivery:

The best way to provide sustained release medication is to place the drug in a delivery system and inject or implant the system into the body tissue. Thermoreversible gels based on poloxamers have been mainly used. ^[77] Poloxamer gel alone or with the addition of hydroxypropylmethylcellulose (HPMC), sodium carboxymethylcellulose (CMC) or dextran was reported for epidural administration of drugs in vitro. ^[78] Pluronic F127 gels incorporating either insulin or insulin-PLGA nanoparticles were evaluated as a successful controlled delivery system. ^[79] Likewise, Intramuscular and subcutaneous administration of human growth hormone using poloxamer gels was tested long acting single dose injection of lidocaine was developed using same gels. ^[80] Noyer et al. invented a new class of injectable controlled release depots of protein which consisted of blends of Pluronics with poly(D, L-lactide)/1-methyl-2-pyrrolidone solutions. ^[81] Thermosensitive hydrogels may also be used for parenteral administration e.g. ReGel® (triblock copolymer PLGA-PEG-PLGA) has been used as a drug delivery carrier for the continuous release of human insulin. Steady amounts of insulin secretion from the ReGel® formulations up to day 15 of the subcutaneous injections were achieved. ^[82] Biodegradable poly(ethylene oxide) and poly(L-lactic acid) hydrogels were developed and exists in their sol-gel transition was reported as a result of temperature change form of sol at an elevated temperature (around 45°C) and forms a gel after subcutaneous injection and subsequent rapid cooling to body temperature. ^[83] Chenite et al. developed novel thermally sensitive combinations of chitosan/polyol salts, which turn into gel implants, when injected in vivo and proved to be a prototype for a new family of thermosetting gels highly compatible with biological compounds. ^[84] Hydrogels formed by xyloglucan were also evaluated as a sustained release vehicle for the intraperitoneal administration of mitomycin. ^[85] PAA/polymethacrylic acid forms a pH-sensitive complex with PEG in situ exhibiting the potential to release drug substances subcutaneously over a period of few days. Alternatively, an aqueous solution containing methyl cellulose (MC) in combination with polymethacrylate yields a reversible gel due to the change of temperature and pH shortly after parenteral administration. ^[86]

(b) Ocular Delivery:

Effective ocular delivery of drug is mostly based on an increase of ocular residence time of dosage form. Ophthalmic hydrogels with enhanced viscosity and mucoadhesive properties is best approach. Among these polymers, in situ gels are preferred since they are conveniently dropped in the eye as a solution, where undergo transition into a gel. Thermosensitive, specific ion sensitive or pH- sensitive hydrogels have been examined for their potential as vehicles for ocular drugs. Poloxamers as thermogelling polymers^[87] have been effectively employed alone or in combination of other such as e.g. PEG ^[88], PAA ^[89], MC, HPMC, CMC ^[90] is often necessary. Ion-sensitive polymers belong to the mainly used in situ gelling materials for ocular drug delivery. Slightly viscous gellan gum solutions in low concentrations (<1%) showed markedly increase in apparent viscosity, when introduced into presence of a physiological level of cations, without requiring more ions than 10–25% of those in tear fluid. The precorneal contact times for drugs can thus be extended up to 20h. ^{[91, 92]} Gellan containing formulations of pilocarpine HCl allowed reduction of drug concentration from 2% to 0.5% obtaining the same bioavailability. ^[93] Alginic acid was exploited for ocular delivery based using ability of gel formation at physiological Ca²⁺ levels. This polymer significantly enhanced the duration of the pressure reducing effect of pilocarpine to 10h ^[94] and carteolol to 8h ^[95] allowing only once a day administration in case of carteolol. Aqueous solutions of PAA that transform into gels upon increase in pH may be used as in situ gelling ophthalmic drug delivery systems. However, the amount of PAA required to form stiff gel upon instillation in the eye is not easily neutralized by the buffering action of tear fluid. Combination of PAA with a suitable viscosity enhancing polymer e. g. HPMC ^[96] or MC ^[97] allows a reduction in the PAA concentration without comprising the in situ gelling properties. The formulation containing Carbopol®940 and Methocel E50LV (HPMC) afforded sustained release of ofloxacin over an 8h period. ^[96]

(c) Peroral Drug Delivery:

Site-specific delivery of drugs to specific regions of the GI tract utilizes pH sensitive hydrogels. Silicone microspheres were prepared using hydrogels made of varying proportions of PAA derivatives and cross-linked PEG which released prednisolone in the gastric medium or showed gastroprotective property. ^[98] Cross-linked dextran, amidaded pectins, guar gum and inulin were investigated to develop a potential colon-specific drug delivery system as potential of faster swelling under high pH. ^[99] Formulations of gellan and sodium alginate for oral delivery of PCM was developed containing complexed calcium ions that undergo gelation by release of these ions in the acidic environment of the stomach. ^[100]

(d) Rectal Delivery:

The rectal route may be used to deliver many types of drugs that are formulated as liquid, semi-solid (ointments, creams and foams) and solid dosage forms (suppositories). Conventional suppositories often cause discomfort during insertion and these also unable to be sufficiently retained at a specific position in the rectum, sometimes they can migrate upwards to the colon, that makes them possible for drug to undergo the first-pass effect. Choi et al. developed novel in situ gelling liquid suppositories with gelation temperature at 30–36°C, using Poloxamer 407 and/ or poloxamer 188 as temperature sensitive polymeric carriers. Bioadhesive polymers were used to modulate the gel strength and the bioadhesive force. ^[101] Charrueau et al. proposed 18% poloxamer 407 solution as a vehicle for short-chain fatty acid enemas. After gelation at 37°C, it allows control release of short chain fatty acids. ^[102] The potential use of thermoreversible xyloglucan gels for rectal drug delivery was investigated and more sustained release of indomethacin was achieved in vitro in comparison with commercial suppositories. ^[103]

(e) Vaginal Delivery:

The vagina, in addition to being an important organ of reproductive tract, serves as a potential route for drug administration. Formulations based on a thermoplastic graft copolymer that undergo in situ gelation have been developed to provide the prolonged release of active ingredients such as nonoxynol-9, progestins, estrogens, peptides and proteins. ^[104] Chang et al. have recently reported a mucoadhesive thermosensitive gel (combination of poloxamers and polycarbophil) which exhibited increased and prolonged antifungal activity of clotrimazole in comparison with conventional PEG-based formulation. ^[105]

(f) Dermal and Transdermal Delivery:

For percutaneous administration thermally reversible gel of Pluronic F127 was evaluated as vehicle for sustained delivery of indomethacin.^[106] In vivo studies suggest that 20% w/w aqueous gel may be of practical use as a base for topical administration of the drug. Poloxamer 407 gel was found suitable for transdermal delivery of insulin. The combination of chemical enhancers and iontophoresis resulted in synergistic enhancement of insulin permeation.^[107]

(g) Nasal Delivery:

Nasal formulations of AEA with chlorpheniramine maleate and tetrahydrozoline hydrochloride were investigated. The findings suggest that liquid AEA formulations facilitate the instillation into the nose and the hydrogel formed on the mucous membrane provide controlled drug release. ^[108]

CONCLUSION:

The number of products based on new drug delivery systems has significantly increased in the past few years, and this growth is expected to continue in the near future. Incorporating an existing medicine into a new drug delivery system can significantly improve its performance in terms of efficacy, safety and improved patient compliance. Delivery systems with a pulsatile or triggered release pattern are receiving increasing interest for the development of various drugs, where conventional systems with a continuous release are not ideal. Some of the current programmable drug delivery systems that employ polymer-based include Covera-HS, Verelan PM, Cardizem LA, Innopran XL, Uniphyl, and naproxen sodium from Andrx Pharmaceuticals. The major drawbacks in these polymer-based triggered release systems arise from biological variations among individuals. The medical and pharmaceutical scientists should now focus concisely over the importance of triggered release of drugs. The key considerations in the design of these systems are their biocompatibility and the toxicity of the polymer-based devices, response to external stimuli, ability to maintain the desired levels of drugs in serum, shelf life and reproducibility.

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Received on 30.12.2010 Modified on 21.01.2011

Research J. Pharm. and Tech. 4(5): May 2011; Page 691-703

Disease	Chronological behavior and symptoms	Drugs used
Arthritis	Increased conc. of c-reactive protein and interleukin-6 in blood. Pain in early morning compared to day time.	NSAIDS, Glucocorticoids
Diabetes mellitus	Increase in blood sugar level after meal.	Sulfonylurea, Insulin
Cardiovascular diseases	BP lowers during sleep cycle but rises steeply at early morning hours.	Nitroglycerin, Calcium channel blockers
Peptic ulcer	Acid secretion increases in night as well as in afternoon.	Proton pump blockers
Asthma	Precipitation of attacks after midnight or at early morning hours.	Beta-agonists, Antihistaminics
Attention deficit syndrome	Increase in DOPA level in afternoon.	Methylphenidate
Hypercholestemia	Increased cholesterol synthesis during night than day time.	HMG CoA Reductase inhibitors

Stimulus	Types of hydrogels	Type Release Mechanism
Thermal	Thermo-responsive hydrogel	Change in temperature changes the polymer–polymer and water–polymer interactions which effect swelling and cause release of drug
Chemical	Hydrogel containing electron-accepting groups	Electron-donating compounds results in formation of charge-transfer complexes effect swelling and release of drug
pH	Acidic or basic hydrogel	Change in pH cause swelling effect in release of drug
Electrical	Polyelectrolyte hydrogel	Applied electric field changes membrane charging and cause electrophoresis of charged drug which effect swelling and release of drug
Magnetic	Magnetic particles dispersed in microspheres	Applied magnetic field changes in pores in gel cause change in swelling which help in release of drug