Controlled release may be defined as a technique or method in which active chemicals are made available to a specified target at a rate and duration designed to accomplish an intended effect. A definition perhaps more acceptable to the chemist and engineer may be: controlled release is the permeation‑moderated transfer of an active material from a reservoir to a target surface to maintain a predetermined concentration or emission level for a specified period of time [1].

Nature and ingenuity have provided a number of well‑known examples of permeation‑controlled processes. Some well‑known natural processes depending on permeation include respiration, osmosis, and the "bloom" or "patina" on grapes and other fruit. Tables 1 and 2 summarize [1], respectively, examples of controlled release found in nature and some invented by man.

Table 1

Controlled Release‑Examples Found in Nature

One‑celled animals controlling the flow of food and waste materials across the exterior membrane

Activation of bacteria spores after long periods of inactivity by exposure to correct environmental conditions

Chameleon's control of skin color

Honey in beehive‑released by eating container

Natural systemic insecticides, e.g., pyrethrum

Creation of an adhesive spider web

Nitrogen fixation in soil by legumes

Insecticide (poison) injection by wasps

Control release of "ink" by squid

Protective odor release by skunk

Ovulation cycle in animals

Controlled decomposition of wood and leaves to produce humus

Controlled release of aroma and nectar at specific times by flowering plants

Natural biological equilibrium reactions

Temperature activated control of blood vessels in skin (blushing)

Pickling, fermentation, etc. processes activated by bacteria

Plants emitting odor to attract or drive off insects and animals (pitcher plant, skunk cabbage)

Barnacle adhesive

Table 2

Controlled Release‑Examples Invented by Man

Artificial kidney utilizing microcapsules of charcoal to adsorb uremic waste products

Detergents packaged in hot‑water‑soluble films

Membrane blood oxygenators

Microcapsular dry food flavors released during mastication

Enteric‑coated capsular pharmaceuticals released by pH differential

Controlled release systemic insecticides absorbed by plants

Glue sticks containing microcapsular adhesives released by pressure

Controlled release foods tailored to the specific needs of the plant

Insecticide microcapsular bait for control of fire ants

Dye markers that release the dye on water contact

Microencapsulated mercaptan for natural gas identification

Slow‑release steroids for birth control applications

Environmentally eroded coatings

Dry flavor powders released on water contact or under heat

Microcapsular injectables possessing controlled release in body fluids

Microcapsular adhesives employing temperature release

Controlled release coatings activated by bacterial attack

Antifouling controlled release coatings

Controlled release insecticide strips

Microcapsular primers and adhesion promoters

Seeds encapsulated in polyvinyl alcohol tapes

I.1 HISTORICAL PERSPECTIVES OF CONTROLLED RELEASE SYSTEMS

The potential benefits of drug therapy were generally not known until the late 1500s when the Swiss physician, Philippus Paracelsus, pioneered the initial use of minerals as drug therapies. Significant advances in drug therapy, however, did not occur until the mid‑1800s. During this period, scientists discovered bacteria and their role in the disease process, developed many new drug therapies, and began to understand how drugs worked in the human body. Almost all of the currently available drug products, however, were unknown to mankind before the 1900s. The era of drug discovery, in the late 1800s and early 1900s, is referred to as the "drug revolution." During this time, many of the newly discovered drug products were developed and used in the treatment of various disease states, as powders or liquids, by oral administration or external application. Very little was known about dosage forms, drug delivery systems, plasma drug levels, in vein administration, or their importance in drug therapy until the early to mid‑1900s.

Since the initiation of pharmacological therapy, maintaining steady therapeutic drug concentration levels, in vivo, has been a major problem. When using intermittent in vein or oral drug administration, the potential disadvantages of such drug therapies include: high plasma concentrations of drugs that may lead to toxicity or low drug levels that cause to sub therapeutic blood levels, and, potentially, cause drug resistance in some instances. In the past, the only way to eliminate the peak and trough plasma levels of drug therapy was to continuously in vein infuse a patient at a constant rate based on the pharmacokinetics of the drug. This type of therapy, however, required constant monitoring of the plasma concentration of the drug by health care professionals and, thus usually cannot be performed at home.

In order to alleviate this problem, a new system for obtaining controlled drug delivery was essential. Research began, in the late 1930s by R. Deansby and A.S. Parkes, on sustained release implantable drug delivery systems [2a]. In 1937, Parks and Dansby presented a paper describing the effects of various hormone preparations on the growth of livestock [2b]. Apparently, they had formulated compressed pure crystalline estrogen pellets and administered them, subcutaneously, to livestock. The results showed a continuous release of the hormone over 3 months in several animals. The idea of hormonal implantation in animals became standard practice in the 1950s and has been shown to increase the growth and feed efficiency of cattle. This discovery sparked an interest in the area of implants leading to further studies and discoveries that have continued to the present time. [2]

I.2 FUNDAMENTAL REQUIREMENTS FOR CONTROLLED RELEASE SYSTEMS

There are several major factors to consider during the development of an implantable drug delivery system. Biocompatibility with the human environment is essential since the system is to be implanted. For a substance to be biocompatible, it must fulfill certain requirements. All agents must be chemically inert, noncarcinogenic, hypoallergenic, and mechanically stable at the implant site. Also, the material should not be physically or chemically modified by local tissue, and the implant should not cause any inflammatory response at the site of implantation. The overall development of these agents is both a time‑consuming and complex process, which consists of many different tests for stability and biocompatibility. If proper biocompatibility is not achieved, many untoward effects can occur such as capsular contracture, unexpected release of the drug, platelet adhesion, tissue damage, or infection of the area surrounding the implant. [2]

I.3 ADVANTAGES OF CONTROLLED RELEASE SYSTEMS

The pharmaceutical literature has shown that the effectiveness of a pharmacological agent can he altered by the route or type of dosage form administered. The more selective a drug is to its site of action, the less the drug needed to be administered. Implantable systems are being developed to keep this targeting prospective in mind. Site‑specific‑targeted drug delivery lowers the dose necessary to be administered, thereby, minimizing side effects. Greater therapeutic efficacy of the drug can also be achieved by such delivery systems as compared to the conventional oral or in vein formulations. The improved sustained release action of these delivery systems also offers additional advantages, such as better patient compliance. Conventional oral therapies may require administration of the dosage form either one, two, or multiple times a day as shown in Figure 1. Some of the implantable systems, on the other hand, have been developed to last as long as 5 years and require only minimal monitoring. Other potential benefits of this type of delivery system are protection of drugs from rapid in vivo metabolism and immediate termination of drug therapy in the case of emergency or toxicity, thus limiting the potential for antiphylactic‑type reactions. [2]

time of day

Figure 1

I.4 LIMITATIONS OF CONTROLLED RELEASE SYSTEMS

Some of the disadvantages of controlled release or the areas that require a thorough appraisal include:

- Cost of controlled release preparation and processing, which may be substantially higher than the cost of standard formulations

- Fate of the polymer matrix and its effect on the environment

- Fate of the polymer additives, such as plasticizers, stabilizers, antioxidants, filler

- Environmental impact of the polymer degradation products following heat, hydrolysis, oxidation, solar radiation and biological degradation

- Cost, time and probability of success in securing government registration of the product. [1]

II. CONTROLLED RELEASE SYSTEMS

In order to overcome drawbacks of systemic delivery mentioned above, a variety of carrier systems for controlled release have been developed [3]. Since appropriate drug release from dosage forms is of critical importance in realizing their therapeutic efficacy, there has been growing interest in developing rate or time controlled oral preparations. To design more advanced dosage forms, various kinds of carrier materials are being developed to deliver the necessary amount of drug, to the targeted site for a necessary period of time, both efficiently and precisely. In one of the examples of controlled release in oral drug delivery systems, Hirayama et al. outline how well the cyclodextrins (a schematic representation of b-cyclodextrin is shown in Fig. 2) satisfied the requirements for a drug carrier, being able to control the rate and time of drug release of water soluble drugs, enhance the drug absorption across biological barriers and to deliver a drug to a targeted site. [4]

Figure 2

The administration other than oral preparations can be named as implantable drug delivery systems. Historically, implantable drug delivery systems have been classified into two major classes: drug implants and implantable pumps containing the drug. The first major class utilizes various types of polymers and polymeric membranes to control the release kinetics of drugs from the delivery systems. This first group of implants can be further subdivided into two different classes: biodegradable and nonbiodegradable systems. The second major class consists of mechanical pump‑type implants, which utilize an infusion pump‑type action to control release of the drug. Due to continual technological advances in this area, a third atypical implant class has emerged. Unique delivery systems such as sustained‑release intraocular systems for the treatment of glaucoma, hydroxyapatite cement systems used in osteomyelitis, and transurethral injection systems for impotence are a few examples of the third group of implants. [2]

II.1 POLYMERIC DRUG DELIVERY

The production of drug‑loaded polymeric pellets and microspheres introduced a new concept in drug administration: Drugs can be delivered to tissues in a sustained, continuous and predictable fashion using polymers as delivery vehicles. Since the discovery of the first controlled‑release polymer systems in the 1960s, new drug delivery systems have become available for clinical use, including steroid‑releasing reservoirs for contraception (Noirplantg and Progestasertg), pilocarpine‑releasing devices for glaucoma therapy (Ocusertg) and a host of new delivery systems for the treatment of cancer. [5]

Drug delivery systems are based on biocompatible polymers, a subset of polymer materials with sufficient biocompatibility and appropriate physical properties to provide controlled delivery. A delivery system, for a specific agent is usually produced by selecting a polymer with the correct physical properties, and designing a composite delivery system that provides the desired rate and pattern of drug release, these designs exploit various mechanisms of drug transport through polymer materials. (Table 3)

Table 3
Examples of polymers used in polymeric drug delivery
Polymers	Class of polymers	Examples
Natural. biodegradable	Proteins	Albumin
	Polysaccharides	Cellulose
		Chondroitin sulfate
		Starch
Synthetic, biodegradable	Polyesters	Poly (lactic acid)
		Poly (lactic-co-glycolic acid)
	Polyanhydrides	Poly [bis (p-carboxyphenoxy)- prop-ane-co-sebacic acid]
		Poly (fatty acid dimer-co-sebacic acid)
	Poly(ortho esters)
Synthetic, non-degradable	Silicone elastomers
	Poly(ethylene-co-vinyl acetate)
	Polyacrylates	Poly isobutylcyanoacrylate
		Poly isohexylcyanoacrylate
		Poly (methyl methacrylate)

II.1.1 Non-Biodegradable Polymers

The first polymeric controlled‑release devices were based on non‑degradable polymers, principally silicone elastomers. In 1964, researchers recognized that certain dye molecules could penetrate through the walls of silicone tubing [5], an observation that lead to the development of reservoir drug delivery systems, which are hollow polymer tubes filled with a drug suspension (Fig. 3).

Figure 3

The drug is released by dissolution into the polymer and then diffusion through the polymer wall, a mechanism that works for any agent that can dissolve and diffuse through either silicone or poly (ethylene‑co‑vinyl acetate), the two most commonly used non-degradable polymers. The Norplant@ 5‑year contraceptive delivery system, approved for use by women in the United States since 1990, is based on this technology. Anticancer drugs can also be delivered by this same approach, as in the case of carmustine release from silicone‑encased drug reservoirs, which provides reasonably constant release for ‑ 30 h. [5]

Solid matrices of non‑degradable polymers can also be used for long‑term drug release (Fig. 4). In comparison to reservoir systems, these devices are simpler (since they are homogeneous and, hence, easier to produce) and potentially safer (since a mechanical defect in a reservoir device, but not a matrix, can lead to dose dumping). On the other hand, it is more difficult to achieve constant rates of drug release with non‑degradable matrix systems; for example, the rate of release of carmustine from a poly (ethylene‑co‑vinyl acetate), matrix device drops continuously during incubation in buffered water [5]. Constant release can sometimes be achieved by adding rate-limiting membranes to homogeneous matrices, yielding devices in which a core of polymer/drug matrix serves as the reservoir. In other cases, water‑soluble, cross-linked polymers can be used as matrices [5]; release is then activated by swelling of the polymer matrix after exposure to water.

Figure 4

Another type of nondegradable system is the magnetically controlled release system. In this type of formulation, small magnetic beads are uniformly dispersed within a polymer (Fig. 5). When the unit is exposed to a biological system, normal diffusion of the drug due to a concentration gradient is seen. However, upon exposure to an external oscillating magnetic field, larger quantities of drug can be released quickly [2]. The major advantage of this type of drug delivery system is the possibility of manipulating the release kinetics of the drug by using external magnetic stimuli.

Figure 5

II.1.2 Biodegradable Polymers

Biodegradable controlled‑release systems offer the advantage of gradual biological elimination without a residual implant structure remaining. Biodegradable polymers, also called bioerodible or bioabsorbable polymers, are synthetic or natural polymers, which hydrolyze in vivo. A large selection of biodegradable polymers, are available as carriers for local drug delivery. Examples of these polymers are shown in Table 4. These polymers degrade in the body at various time periods, from a few days up to 2‑3 years. These polymers degrade into non-toxic acids or alcohols that are readily excreted from the body. The in vivo elimination time is determined by, the nature of the polymer chemical linkage, the solubility of the degradation products, the size, shape and density of the device, the drug and additive content, the molecular weight of the polymer, and the implantation site. Many of the site‑specific applications of drugs are for periods of several weeks, requiring polymer carriers that degrade and are eliminated from the body soon after. It is believed that for many of the applications the polymer should be eliminated within 6 months after implantation. From the available polymers, polyanhydrides, collagen and copolymers of lactide and glycolide are useful for short‑term drug release and device elimination in vivo [5]. There are three general mechanisms of polymer hydrolysis:

1. Erosion of cross‑linked polymers with hydrolytically unstable cross‑links. Polymer chains are freed from the bulk matrix as the cross‑links are hydrolyzed. This mechanism is useful for drugs with low water‑solubility or large macromolecules.

2. Solubilization of water‑insoluble polymers by hydrolysis, ionization or protonation of a side group, without any significant change in polymer molecular weight. This mechanism is mainly used in topical or oral applications.

3. Erosion of water‑insoluble polymers with labile backbone bonds. Hydrolysis of the polymer backbone produces low molecular weight, water‑soluble molecules. This mechanism is most useful for systemic administration of drugs from subcutaneous, intramuscular or intraperitoneal implantation sites. [5]

Table 4

Common biodegradable polymer carriers
Polymer	Polymer linkage	Principal degradation product	Elimination time^a (months)
Poly(lactic acid)	-CO-O-	Lactic acid	12-24
Poly(lactic-co -glycolic acid)	-CO-O-	Lactic and glycolic acid	6-12
Poly(glycolic acid)	-CO-O-	Glycolic acid	2-4
Poly(caprolactone)	-CO-O-	Hydroxypentanoic acid	18-24
Poly(hydroxybutyrate)	-CO-O-	Hydroxybutyric acid	18-24
Poly(orthoester)	-CO-O-	Alcohols	12-24
Poly(alkane anhydride)	-CO-O-CO-	Aliphatic diacids	0.2-4
Gelatin collagen	-CO-NH-	Amino acids	0.2-1
Oxidized cellulose	-C-O-CH-O-	Alcohols, CO₂	0.2-1
Poly(phosphazene)	-N=P-	Phosphates, ammonia	6-18
^aElimination times vary depending on implant size and shape, density, implantation site and molecular weight.

Bioerodible polymers can be reservoir or matrix devices. In a reservoir system, a core of drug is surrounded by a polymer and the rate‑limiting step in drug release is the diffusion of drug outward through the polymer. This rate can be controlled or adjusted by changing the nature of the bioerodible membrane. In a matrix system, drug is dispersed uniformly throughout a solid polymer. Drug release is a product of either diffusion through or erosion of the polymer. In order to he clinically useful, biodegradable polymers must also be biocompatible. The polymer and its degradation products must either be safely eliminated from the body or metabolized to non-toxic substances. [5]

There are many natural and synthetic polymers that have potential application as biodegradable drug delivery systems. These include polyamides, polyesters, polyanhydrides, poly(orthoesters), and various polysaccharides. [5]

II.1.3 Liposomes

Although many of the materials used to prepare liposomes are not polymers, these amphiphilic materials do form very ordered structures with specific orientations similar to that of polymers. More recently, polymers have been used to stabilize liposomes either by a coating of the liposome or using materials, which can polymerise after the liposome has been formed [6].

A revival of interest in liposomes for drug delivery occurred in the last few years as a result of the design of so called “stealth” liposomes, which are capable of withstanding long-term circulation in the blood stream. Liposomes with encapsulated magnetic particles (magnetoliposomes) have been suggested for magnetically controlled, targeted chemotherapy [3].

II.2 IMPLANTABLE PUMPS

Many different drugs require external control of delivery rate and volume. Such control cannot be obtained when using biodegradable or nondegradable delivery systems with the exception of the magnetic type delivery systems. Pump systems have been used to provide the control needed in these situations. Recently, due to the availability of advanced microtechnology, it has been possible to create pump systems small enough to implant, subcutaneously, for drug delivery. This allows the patient to maintain the control of drug release without the need for an external pump system.

Pump systems differ from other implantable systems due to their mechanism of drug delivery. Pump systems release drugs through a pressure difference generated gradient that results in the bulk flow of a drug at controllable rates [2]. To date, five different types of implantable pump systems have been tested including infusion pumps, peristaltic pumps, osmotic pumps, positive displacement pumps, and controlled release micropumps.

Infusion Pumps

Infusion pumps are implantable mechanical systems that utilize a fluorocarbon propellant to administer the drug, in vivo. Such pumps were initially developed for the administration of insulin to diabetic patients. Infusaid (Infusaid Corp. Sharon, MA) was one of the first commercially available pumps for this use. Normally, insulin‑dependent diabetics require injections once or twice daily. This type of dosing results in abnormal peaks and valleys in blood glucose levels. It is believed that such poor control of blood glucose levels may lead to diabetic complication such as heart and kidney disease [2]. It is felt that continuous insulin infusion using such pumps may help eliminate these risk factors in the diabetic population.

The pump consists of a disc‑shaped cannister made of light‑weight biocompatible titanium, which contains a collapsible welded bellow [2]. The bellow separates the canister interior into two separate chambers. The first chamber contains the fluorocarbon propellant and the second contains the insulin formulation (Fig. 6). The gas pushes the drug through a filter and also a flow regulator that provides a constant rate of drug administration at a given temperature. The delivery rate is adjusted by changing the drug concentration in the pump reservoir [2]. The advantage of this system involves the fact that no external energy source is needed to drive the pump action. When the pump reservoir must be refilled, an injection of drug through a membrane consisting of a self‑sealing silicone rubber and Teflon septum is administered. The force of the injection recompresses the fluorocarbon propellant thereby recharging the system. In addition to insulin therapy, the use of this pump system in the delivery of anticoagulant and chemotherapeutic agents has also been investigated [2].

Peristaltic Pumps

Peristaltic pumps consist of rotary solenoid‑driven systems that run via an external power source, which is usually a battery [2]. Peristaltic systems, like the infusion pump systems, are filled through a silicone rubber septum, and can be used for several years depending on the life span of the battery‑powered system (Fig. 7). The advantage of this type of system is that the rate of drug administration can be controlled by an external remote control system. These systems, however, have proven to be very costly, and, thus, have not been seen in standard practice to date [2].

Figure 6

Figure 7

Osmotic Pumps

Osmotic pumps have proven to be the most popular type of implantable drug delivery systems. The osmotic pump, also known as Oros or the gastrointestinal therapeutic system, was first described by Theeuwes and Yum, and released for use by Alza Corporation [2]. This pump consists of a drug reservoir surrounded by a semipermeable membrane. The surrounding membrane allows a steady influx of water and biological fluid into the reservoir through the process of osmosis. The hydrostatic pressure build‑up from this influx causes a steady release of the drug from an opening in the membrane called the drug portal. The rate of drug release is constant or zero‑order until the drug within the reservoir is completely depleted [2]. Changing the rate of drug administration of these systems can only occur by changing the structure of the semipermeable membrane, which requires removal of the system [2].

Osmotic pump systems containing hydromorphone have been subcutaneously implanted for the use of pain management. Results have shown that Aizet's osmotic pumps release 262 mg/h of hydromorphone to produce stable plasma concentrations of approximately 30‑40 mg/ml over a 2‑week period [2]. This type of delivery system is advantageous over other systems since the "initial burst effect," seen in other forms of degradable or nondegradable matrix systems, does not occur [2]. The prolonged release of drug at a constant rate has been shown to be effective in the treatment and management of chronic pain. Therefore, such systems may be used more extensively in the future.

Positive Displacement Pumps

Positive displacement pumps have been developed to provide continual insulin delivery in diabetic patients. Most of these systems utilize piezoelectric disk benders affixed to flexible tubing [2]. Such pumps are made by first exposing the disks to certain voltages so that they form spherical surfaces. The bellow‑type system is then connected to a drug reservoir via a three‑way solenoid driven valve. When exposed to electrical pulses, the valves in the pump open or close depending on the direction of the pulse. This action causes the release of drug in a controlled manner based on the rate of the electrical pulse. Other types of positive displacement pumps using similar designs are currently being developed for the delivery of insulin [2].

Controlled Release Micropumps

Controlled release micropumps have been developed using diffusion across a rate‑controlling membrane to affect basal delivery, while a rapidly oscillating piston acting on a compressible disc of foam increases the rate of drug delivery [2]. This type of pump system dose, not require an external power source. Suitable power to run this pump is achieved via the concentration difference between the drug reservoir and the site of action. This concentration difference causes the diffusion or basal delivery of the drug. To enhance the delivery of the drug, a current is run across the solenoid coil. This current causes compression of the foam disc via a steel piston. The pressure difference, achieved by moving the steel piston, causes enough force to excel delivery of the drug. This pump is currently being developed for use in insulin delivery. Studies have shown that such pumps can administer average daily flow rates necessary for adequate insulin therapy using commercially available insulin solutions of 100 U/ml [2]. One particular unit studied fulfilled the requirements of basal insulin dosage and could be controlled over a wide range of basal supply needs. Preliminary results showed that, if the device was implanted, depending on the endurance of the foam disc, it could be used up to 1 year [2].

III. RECENT APPLICATIONS

III.1 CONTROLLED RELEASE IN CANCER THERAPHY

While the search for new anticancer agents is in progress, optimization of delivery for existing drugs will remarkably improve the current scenario in the management of cancer. When anticancer agents are delivered systemically, they must cross several barriers to reach their site of action within tumor cells [5]. Drug molecules must permeate through capillary walls, diffuse through the extracellular space, traverse the tumor cell membrane and reach the proper intracellular target. Unstable agents may be subject to spontaneous or metabolic elimination during this process; consequently, drug concentrations at the target may be low. Higher systemic doses could provide additional antitumor activity at the target, but may also result in undesirable toxic effects. Polymeric delivery systems can be used to provide better or safer chemotherapy by either prolonging the duration of systemic therapy or focusing drug therapy in a particular tissue region.

Polymeric microparticles for sustained chemotherapy

Conventional treatments for hormone‑responsive cancers, such as carcinomas of the breast, prostate and endometrium, require frequent injections of hormone agonists [5]. For example, injections of synthetic agonistic analogs of luteinizing hormone-releasing hormone inhibit tumor growth in the prostate by lowering plasma testosterone levels [5]. Since the plasma half‑life of luteinizing hormone-releasing hormone analogs in humans is short (e.g. 3.6 h after subcutaneous injection and 2.9 h after intravenous injection [5]), daily injections are necessary to maintain concentrations that are sufficient to suppress testosterone production. Unfortunately, daily injection is inconvenient and often produces an undesirable testosterone release response in patients [5].

To lower the frequency of injections, (luteinizing hormone-releasing hormone analog)‑loaded poly(lactic-co-glycolic acid) microspheres (Lupron Depot^® [5]) and cylindrical implants (Zoladex^® [5]) were produced. Following injection of the microspheres, for example, scrum testosterone levels are reduced for 4 weeks [5]. When the microspheres are injected subcutaneously or intramuscularly, they are immobilized at the injection site. The drug molecules are released from the immobilized spheres, diffuse locally in the tissue, and then enter the vascular system, which distributes the drug throughout the body. Unlike free drug injection, the effective half‑life for therapeutics administered by microspheres is much longer and the total amount of administered agent is reduced. Even longer periods of drug release are potentially achievable using this approach [5].

Prolonged systemic delivery appears particularly useful for hormonal anticancer therapies. For example, medroxyprogesterone acetate is an androgenic steroid often used in the hormonal therapy of breast cancer in women [5]. A single subcutaneous injection of medroxyprogesterone acetate‑loaded poly(lactic-co-glycolic acid) microspheres inhibits both growth and initiation of chemically induced mammary tumor in rats, suggesting that long‑term, low‑dose delivery systems may be useful for the treatment and prevention of breast cancer. [5]

Polymeric implants for sustained interstitial chemotherapy

The most common tumors are solid tumors of the lung, breast, colon and prostate. Generally, these tumors do not respond well to conventional systemic chemotherapy or radiotherapy, especially when the tumors are large or poorly vascularized [5]. Solid tumors of the brain occur less frequently, but are among the most difficult to treat because the blood‑brain barrier prevents the entry of most intravascular agents into the brain. If the tumor is operable surgical removal is the preferred therapy, but many brain tumors, such as malignant glioma, can recur; recurrence is usually within a 2‑cm margin of the excision site [5]. Local chemotherapy, provided directly at the site of tumor resection, is a reasonable approach for preventing recurrence.

Polymer drug delivery systems provide an opportunity to deliver high, localized doses of chemotherapy for a prolonged period after tumor resection. The most extensively characterized local delivery system is carmustine‑loaded poly(p-carboxyphenoxypropane-co-sebacic acid) polymer pellets, which have been tested for biocompatibility, efficacy, dose escalation and pharmacokinetics in animals. [5] Clinical trials in the United States and Europe have demonstrated the safety and efficacy of this approach in humans [5]. Certain chemotherapeutic agents that do not cross the blood-brain barrier, such as carboplatin, 4‑hydroperoxycyclophosphamide and taxol (commercially called paclitaxel), are effective when directly delivered by polymer implants to animals with experimental tumors [5]. This approach appears to be useful for novel chemotherapy agents as well, including camptothecan (a topoisomerase I inhibitor), angiogenesis inhibitors (minocycline and heparin /cortisone) [5].

While direct delivery to solid tumors by implanted polymers has been studied most extensively in the brain, polymeric pellets can potentially deliver chemotherapeutic agents locally to extracranial tumors as well. Polylactide needles were used to deliver 5‑fluorouracil for the treatment of hepatomas in rats; similar materials were used to deliver adriamycin intratumorally in animals with mammary carcinoma. Bleomyein‑loaded poly(lactic acid) cylinders were implanted in the mediastinum of dogs for targeted chemotherapy for esophageal cancer [5]. Cisplatin has been delivered by a variety of polymers: In polyanhydrides, for the local treatment of squarnous cell carcinoma of the head and neck, and also in poly(methyl methacrylate), fibrin glue and poly(ethylene-co-vinyl-acetate), for the local treatment of osteosarcoma [5]. In all of these studies, intratumoral administration of anticancer agents by polymers resulted in higher drug activity in the tumors, when compared to more conventional delivery strategies.

The effectiveness of certain chemotherapeutic agents that do not cross the blood-brain barrier, when directly delivered by polymer implants is mentioned before. Paclitaxel, a new antineoplastic agent, is one such drug deserving attention in the field of regional drug delivery, offering immense pharmacokinetic as well as therapeutic advantage via localized delivery. [7] Drug delivery systems explored so far for localized paclitaxel delivery are microspheres, surgical pastes and implants. (Table 5)

Table 5

Delivery systems developed for localized paclitaxel delivery
Polymeric device	Advantages	Disadvantages
Microspheres	Chemoembolization, no dose dumping	Usually first order release
Surgical pastes	Direct injection	Less consistent performance
Biodegradable implants	Large functional life-time	Possible risk of rejection and dose dumping

Microspheres

Microspheres (0.6% paclitaxel) composed of a blend of biodegradable poly lactic acid and a non‑degradable ethylene vinyl acetate copolymer were prepared by the solvent evaporation method [7]. A blend of 1:1 ethylene vinyl acetate: poly lactic acid was selected as a composite material based on the observation that the microspheres could be easily prepared due to decreased problems of tackiness and coalescence and exhibited sustained drug release with increased functional life‑time. Though the in vitro studies indicated an initial burst effect, the overall release rates were very slow. Therefore, characterization of the crystal form of the drug in the polymer might disclose a plausible reason for the slow release.

Further studies with isopropyl myristate revealed that isopropyl myristate significantly increased the release rate of paclitaxel in vitro from Paclitaxel‑ isopropyl myristate ‑poly(lactic-co‑glycolic acid) microspheres and it was found that the release is governed by diffusion in the matrix. One more study was conducted to examine the significance of molecular weight and copolymer ratio of poly(lactic-co‑glycolic acid) on drug release rate and polymer degradation profiles. The release rates increased with molecular weight and lactic acid content of the poly(lactic-co‑glycolic acid) matrix due to the greater number of micropores on the surface as evident from the scanning electron microscopy scans. Lactic acid being more hydrophobic than glycolic acid must have facilitated drug release. In the presence of isopropyl myristate, polymer with both higher molecular weight of poly(lactic-co‑glycolic acid) and lactic acid content precipitated faster during the preparation by the solvent evaporation method, accounting for the larger number of micropores on the surface. However, poly(lactic-co‑glycolic acid) microspheres (lactic acid:glycolic acid, 75:25, molecular weight = 10 000) gave zero‑order release for 3 weeks probably because of perfect balance between isopropyl myristate mediated diffusion and matrix erosion. [7]

Surgical pastes

Surgical pastes are monolithic drug containing devices prepared by thermal processing of polymeric materials having suitable glass transition temperatures or by in‑situ solvent incompatibility for direct injection at a desired site.

Polycaprolactone pastes of paclitaxel (1‑30% w/w) were prepared and characterized with varying concentration of methoxypolyethyleneglycol (0‑20%). The preformulation studies conducted revealed that methoxypolyethyleneglycol (MW = 350) blended more homogeneously with polycaprolactone matrix (MW = 20 000‑27 000) reducing its melting point sufficiently for injection in the presence of drug as compared to other formulations containing other molecular weight fractions of polyethyleneglycol, a blend of polycaprolactone:methoxypolyethyleneglycol (4:1) melted at 50.4°C. Differential scanning colorimetry studies revealed that both methoxypolyethyleneglycol and paclitaxel decreased the melting point of polycaprolactone matrix but increased its crystallinity, thus increasing its functional life‑time. Though tensile strength of the pastes was decreased by methoxypolyethyleneglycol in a concentration dependent manner, they did not disintegrate, possibly due to the increased crystallinity. Microscopy showed the monolithic solution nature of surgical paste at loadings up to 10%, and at loadings greater than 5% they were monolithic dispersions, suggesting that the release mechanism of the former is different from that of the latter and is loading dependent. In vitro release studies showed that the incorporation of methoxypolyethyleneglycol did not result in increased release rates. Hence, the increased crystallinity as confirmed from differential scanning colorimetry studies retarded release due to decreased molecular diffusion coefficients. [7]

Biodegradable surgical pastes of poly(D,L‑lactic acid)-polyethyleneglycol- poly(D,L‑lactic acid) copolymer and poly(D,L‑lactic acid) + polycaprolactone blends of paclitaxel at 20% loading were fabricated. As described previously, an increase in the polyethyleneglycol content resulted in the crystallinity of the block copolymer, but in vitro release studies indicated that release rates increased with polyethyleneglycol content in contrast to the earlier group. Pastes with 10% polyethyleneglycol had high melting point (> 60'C) unsuitable for injection, whereas those with greater than 401/0 disintegrated due to swelling of the incorporated polyethyleneglycol. This was attributed to the reduced polymer MW due to polymer breakdown and release of polyethyleneglycol as supported by the data from gel permeation chromatography (GPC). In vitro release studies from poly(D,L‑lactic acid) + polycaprolactone blends showed an increase in the release rate with poly(D,L‑lactic acid) content and pastes with poly(D,L‑lactic acid) content greater than 80% disintegrated because poly(D,L‑lactic acid) (MW = 800) degraded rapidly while polycaprolactone served as holding material. The formulations poly(D,L‑lactic acid):polycaprolactone (90:10) and poly(D,L‑lactic acid)-polyethyleneglycol- poly(D,L‑lactic acid) (30% polyethyleneglycol) caused tumor weight regression by 54 and 40%, respectively, in subcutaneously established MDAY‑D2 tumors in mice after local delivery via injection at tumor site. [7]

The problem of increased crystallinity could be overcome by replacing methoxypolyethyleneglycol with coprecipitated microparticles of paclitaxel with various water soluble polymers into the polycaprolactone matrix by melt technique corresponding to a 20% loading of the drug. In vitro release rate profiles followed the square root of time relationship indicating a diffusion related mechanism. Paclitaxel release from the polycaprolactone matrix increased in the presence of additives in the same order as the rate of swelling and with the proportion of additive and the size of drug‑additive microparticles, but this also resulted in the disintegration of the matrix and therefore clinically useless formulations. The additives with a higher swelling nature exerted tumor pressure rupturing the polymer barrier between adjacent particles creating micro‑channels and thus facilitating the escape of drug molecules from the matrix. A similar observation was made from in vivo studies after peri‑tumoral injection of the molten paste with established palpable tumors. Paclitaxel‑gelatin‑polycaprolactone pastes (20:20:60) produced a reduction of 63± 27% in tumor mass. [7]

Biodegradable implants

Drugs that readily cross the blood-brain barrier are more effective against CNS tumors than drugs that do not cross the blood-brain barrier because of their size and lipophilicity. As a generalization, drugs that cross the blood-brain barrier to produce appreciable brain drug levels are either hydrophilic with MW<160 or hydrophobic MW<400. The much hydrophobic paclitaxel molecule (MW = 853.9) has poor blood-brain barrier penetration capacity and hence, the use of implants for intracerebral delivery is justified. [7]

Intracerebral chemotherapy is emerging as an important therapeutic approach in the treatment of recurrent malignant gliomas because of the pharmacokinetic advantage offered bypassing the blood-brain barrier and also due to the fact that malignant gliomas recur within a few centimeters away from the tumor excision site. Drug loaded polymer discs (carmustine, 4‑hydroxycyclophosphamide or paclitaxel) of poly[bis (p‑carboxy phenoxy)propane‑sebacic acid] (20:80) at 20% loading were implanted intracranially in cyanomolgus monkeys. [7] The rate of release of paclitaxel remained constant (~3 µg/day) over a 30-day period of in vitro study with 7% release in 100 days. The blood level of paclitaxel was low in comparison to carmustine and 4‑hydroxycyclophosphomide due its highly hydrophobic nature and low local concentrations attained at the polymer disk‑tissue interface. Pharmacokinetic analysis of drug distribution data confirmed that interstitial fluid convection in addition to concentration gradients contributed to drug transport and distribution in the brain. Similar slow release of paclitaxel, 15% in 100 days, from poly[bis(p‑carboxyphenoxy)propane‑sebacic acid], and 17% in 100 days from poly(fatty acid dimer‑sebacic acid) (1:1) was observed. [7]

However, faster release rate (45‑65% in 30 days) of paclitaxel was observed from relatively more hydrophilic poly[bis(p‑carboxyphenoxy)propane‑sebacic acid] (20‑40% loading) and the implants were found to be promising in an experimental glioma model. The implants doubled to tripled the median survival of rats bearing tumor. Paclitaxel concentrations remain elevated for at least 1 month after implantation in brain and offered a good pharmacokinetic advantage due to prolonged exposure to therapeutic concentrations bypassing the blood-brain barrier. Since paclitaxel is extremely lipophilic, the more the hydrophilicity of the polymeric device faster is the release rate, and the desired release rate can be achieved by judicious selection of the polymeric material. [7]

A recent application of controlled release polymer-based chemotherapy was reported by Langer et al. [8]. To prepare the polymeric compound they selected monomers hypothesized to be non-toxic. In one example they used a hydrophobic monomer, carboxyphenoxypropane, and a slightly less hydrophobic monomer, sebacic acid. To make these polymers, numerous polymer chemistry challenges had to be addressed, such as synthesis of high‑molecular‑weight polyanhydrides. However, by optimizing the time and temperature of the polymerization reaction and using appropriate catalysts, polyanhydrides with molecular weights up to 250000 were synthesized. From an erosion standpoint, using 100% carboxyphenoxypropane and no sebacic acid, making a very hydrophobic polyanhydride, about 8% of the polymer matrix dissolves in 14 weeks. A millimeter‑thick disk will take about 3 or 4 years to dissolve. However, when they added as a co‑monomer 15% sebacic acid, the polymer matrix dissolves faster; with 79% sebacic acid, the polymer matrix completely dissolved in 2 weeks. Thus, by specifying the monomer ratio, these polymers can be made to last for essentially whatever length of time is desired. If drug is uniformly distributed inside such a polymer matrix, generally at 10 wt % or less, it will be released at the same rate at which the polymer dissolves.

To apply this technology to clinical problems, they began developing a better way to treat glioblastoma multiforme, a uniformly fatal form of brain cancer. Untreated, the median life expectancy is 4 weeks. Surgery changes that to 16 weeks; surgery and radiation to 40 weeks; and surgery, radiation, and chemotherapy to 50 weeks. Furthermore, patients often are subjected to an undesirable quality of life in which they may experience repeated surgeries and are given one of the most toxic anticancer drugs, carmustine. Typically taken intravenously, the drug is transported throughout the body, causing damage to the liver, kidneys, and spleen. Our objective was to develop a localized therapy that might provide a new way to treat patients without these side effects.

Their goal was to enable the surgeons to operate, removing as much tumor as possible, and then before closing the patient, adding polymer‑ carmustine disks. They desired a polymer that was biodegradable so that it would not remain in the brain over long time periods, and one that was surface eroding because the drug is toxic. They placed carmustine, which normally has a half‑life of 12 min, in the polymer matrix. Once in the polymer matrix, the drug lasts as long as the polymer does because it is protected from degradation. The surgeons inserted this delivery system in the brain where it was needed and did not cause damage to various organs.

Initially, they experimented with the polymer in vitro assays using mammalian cells to examine toxicity. Then, safety studies were done in rats, rabbits and monkeys. In all cases the polymers were safe. Then human clinical trials began. In these studies, the disks, which are about the size of dimes, were placed into the brains of patients who were being operated on for cancer.

The initial human studies examined safety. Blood tests showed no toxicity. Subsequent studies addressed efficacy. One study showed that after a year, 63% of patients were alive in the treated group and 19% in the control group. After 2 years, 31% of patients4ere alive in the treated group and 6% in the controls. This marks the first time in over 20 years that a new brain cancer treatment was approved.

III.2 OTHER APPLICATIONS OF BIODEGRADABLE POLYMERIC SYSTEMS

A wide variety of natural and synthetic biodegradable polymers have been investigated for drug targeting or prolonged drug release. However, only a few of them are actually biocompatible. Natural biocompatible polymers like bovine serum albumin, human serum albumin, collagen, gelatin, and hemoglobin have been studied for drug delivery [9]. The use of these natural polymers is limited due to their higher costs and questionable purity.

Since the last two decades, synthetic biodegradable polymers have been increasingly used to deliver drugs, since they are free from most of the problems associated with the natural polymers [9]. Poly(amides), poly(amino acids), poly(alkyl‑a‑cyano acrylates), poly(esters), poly(orthoesters), poly(urethanes), and poly(acrylamides) have been used to prepare various drug loaded devices [9]. Amongst them, the thermoplastic aliphatic poly(esters) like poly(lactide), poly(glycolide), and especially poly(lactide-co-glycolide) have generated tremendous interest due to their excellent biocompatibility and biodegradability [9].

Figure 8

Various polymeric devices like microspheres, microcapsules, nanoparticles, pellets, implants, and films have been fabricated using these polymers for the delivery of a variety of drug classes.

Jain et al. discussed the various traditional techniques of preparing various drug loaded poly(lactide-co-glycolide) devices [9]. They have also described a method, which can be terminated at various stages to yield different syringeable mixtures. (Fig. 8) These on injection came into contact with water and formed solid matrix type implant or microspheres (in situ formed implant or microspheres respectively) entrapping cytochrome c. The method can also be made to produce injectable, isolated microspheres. The protein, cytochrome c, was encapsulated by these three devices and was released in a controlled fashion.

Hydrogels are polymeric materials that do not dissolve in water at physiological temperature and pH but swell considerably in aqueous medium [10]. Hydrogels are of special interest in controlled release applications because of their soft tissue biocompatibility, the case with which the drugs are dispersed in the matrix and the high degree of control achieved by selecting the physical and chemical properties of the polymer network. Hydrogels consist of polymer chains cross-linked to each other to create a tangled mesh structure, providing a matrix for the entrapment of drugs. These properties conducted a considerable research on hydrogels and their usage in controlled release technology.

In of the studies Risbud et al. have reported, pH‑sensitive chitosan‑polyvinyl pyrrolidinone hydrogels as a controlled release system that can be used in gastric environment. [10]

In case of diseases involving peptic ulcers, it has been demonstrated that Helicobacter pylori is one of the major causative agents. This bacterium releases the enzyme urease, which converts urea into ammonia and bicarbonate, which aids in neutralising the acidic medium and allows the bacteria to colonise the gastric mucosa. Amoxicillin and metronidazole, which are effective in treating H. pylori under in vitro conditions, score poorly when used to treat infections in an in vivo situation. The failure of these antibiotics has been proposed to be an outcome of sub‑effective bactericidal concentrations available at the site and their instability following oral administration [10].

Cationic hydrogels with pH‑sensitive swelling properties have been proposed previously as candidates for stomach‑targeted drug delivery systems [10]. Such matrices can be used to provide adequate drug release in gastric (low pH) environments. Chitosan, a cationic polysaccharide, is obtained by, alkaline deacetylation of chitin, the principal exoskeletal component in crustaceans. Chitosan is reported to be non‑toxic and bioabsorbable, and has been explored for the release of many drugs. Risbud et al. have reported on a pH‑sensitive chitosan‑polyvinyl pyrrolidinone, semi-interpenetrating polymer network‑based controlled release antibiotic delivery system that is well‑suited for use in a gastric environment. Figure 9 represents the antibiotic release from hydrogels, which was studied in solutions with pH values of 1.0, 2.0, 3.0. [10]

Their process has generated matrices with high porosity (pore diameter 39.20 ± 2.66 µm) that exhibited superior pH‑dependent swelling properties, which could be attributed to their porous nature. The increased swelling of hydrogels under acidic conditions was due to the protonation of the primary amino group on chitosan. With 3 h, membranes released 73.2 and 51% of amoxicillin in solutions of pH 1.0 and 2.0 respectively. So they suggest that their hydrogels could serve as potent candidates for antibiotic delivery in an acidic environment. [10]

Figure 9

In another study Han et al. [11] prepared lactilol-based poly(ether polyol) hydrogels. They outlined that swelling ratio of their hydrogels decreased as the cross‑linking ratio increased. They suggested that, free volume between cross‑linked chains might act as an important factor during swelling to trap free water molecules. They also reported that, the swelling ratio also decreased with increasing temperature, NaCl concentration, and glucose concentration. The pH did not significantly affect the swelling ratio of the hydrogels. The hydrogels lose associated water molecules by temperature increase or when the water is competitively hydrated by salt and glucose. (Fig 10) The swelling ratio of the hydrogels changed reversibly in response to the environmental solution conditions. Han et al. proposed that, this property may be used for developing environmentally responsive delivery systems or biosensor probes.

Figure 10

The aspirin release profiles from the hydrogels were sensitive to the cross‑linking ratio. Highly cross‑linked hydrogels showed slower release of aspirin than the lightly cross‑linked hydrogels. The diffusivities of aspirin from cross‑linked hydrogels were found to decrease as the cross‑linking ratio increased (Fig 11). Therefore, the results suggest that, the release rate of incorporated chemical may be controlled, by changing the cross-linking ratio of the hydrogels. [11]

Figure 11

In another study chemically cross-linked gelatin‑chondroitin sulfate hydrogels were prepared by Kuijpers et al. for the controlled release of small cationic proteins. [12] The amount of chondroitin sulfate in the gelatin gels varied between 0 and 20 wt %. The chemical cross-link density, the degree of swelling, and the theological behavior were determined to characterize the cross‑linked hydrogels. [12]

Chemically cross‑linked gelatin chondroitin sulfate hydrogels were loaded with lysozyme, and the release was measured using phosphate‑ buffered saline. The lysozyme loading capacity of the hydrogels significantly increased with increasing chondroitin sulfate content of the gels. Compared to plain gelatin gels, the release rate of lysozyme slowed for the hydrogels containing 5 and 10 wt % of chondroitin sulfate, while the release was faster for hydrogels containing 20 wt % of chondroitin sulfate. The permeation of lysozyme through gelatin‑ chondroitin sulfate gels was measured using a two‑compartment diffusion cell, and the effective diffusion coefficient was calculated. The effective diffusion of lysozyme in the gels was also qualitatively studied using fluorescence recovery after photobleaching. The Langmuir isotherms of lysozyme adsorption to gelatin‑ chondroitin sulfate gels and the lysozyme diffusion in the gels in the absence of electrostatic interactions were determined to evaluate the contributions of unspecific interaction between lysozyme and chondroitin sulfate and diffusion to the release. Both the interaction and the diffusion increase with increasing chondroitin sulfate content of the hydrogels, which resulted in a minimum value of the effective release rate for gels containing 5 wt chondroitin sulfate. [12]

Another recent study is done by Kunou et al. to treat cytomegalovirus retinitis by long-term sustained release of ganciclovir from a biodegradable scleral implant. [13]

Cytomegalovirus retinitis is the most common opportunistic ocular infection and the major cause of visual loss in patients with acquired immune deficiency syndrome (AIDS). Cytomegalovirus retinitis occurs 20 to 25% of patients with AIDS during the course of their illness. With the recent advent of highly active antiretroviral therapy, the incidence of cytomegalovirus retinitis has been reported to decrease significantly and the control has become easier because of immune recovery. In some patients, however, cytomegalovirus retinitis remains a significant problem, as patients who failed in the highly active antiretroviral therapy developed new or reactivated cytomegalovirus retinitis. Therefore, treatment modalities are still required for patients who do not undergo or do not respond to highly active antiretroviral therapy. Currently, there are six treatments for cytomegalovirus retinitis: intravenous and oral ganciclovir, intravenous foscarnet intravenous cidofovir, and ganciclovir intravitreal implant and intravitreal fornivirsen. Administration of intravenous ganciclovir, foscarnet, and cidofovir is effective at slowing the progression of cytomegalovirus retinitis. However, long-term treatment and high‑dose intravenous injections have caused side‑effects: in the case of ganciclovir, dose related bone marrow suppression and neutropenia have been reported. Intravitreal injections of ganciclovir were applied to the patients needing a higher level of the drug within eye. The injections had to be repeated at least every one or two weeks and might be associated with a risk of cataract, retinal detachment, and endophthalmitis. [13]

Recently, a ganciclovir intravitreal implant has been developed as a local treatment option that avoids systemic side‑effects and does not involve intravitreal injections. This implant is made of ethylene‑vinyl acetate copolymer and poly(vinyl alcohol), and releases ganciclovir by passive diffusion through a small opening in ethylene‑vinyl acetate copolymer at the base of the device for 6‑8 months. This device is not biodegradable and requires a surgical procedure for its removal or repeated implantation when drug release is completed, producing retinal detachment, endophthalmitis, vitreous hemorrhage, and others [13].

Poly(lactide‑co‑glycolide) has been mentioned before as most widely studied biodegradable polymer because of the long history of its safe clinical use as resorbable sutures. However, drug release kinetics from poly(lactide‑co‑glycolide) matrices typically exhibit tri‑modal behavior starting with a high release rate stage attributed to the release of surface‑deposited drug, followed by a slow release stage attributed to diffusion through the matrix, and a final rapid release stage attributed to dumping of the residual drug during the later stage of degradation. Drug release kinetics from bulk‑eroding poly(lactide‑co‑glycolide) matrices is complex because the polymer phase properties change continuously during degradation, resulting, in drastic changes of drug, diffusivity and permeability. [13]

Kunou et al. developed a poly(lactide‑co‑glycolide) scleral device and implanted it at the pars plana to release ganciclovir directly into the vitreous without disturbing the transparency of the ocular medium. In previous studies using pigmented rabbits, the ganciclovir concentration in the vitreous after implantation of ganciclovir‑loaded scleral implants was maintained within the therapeutic range for cytomegalovirus retinitis for 3 months. The previous implant, however, had sonic disadvantages such as the second burst in the late phase of release.

In this study, the ganciclovir release rate from scleral implants was modified by blending high and low molecular weight poly(D,L‑lactide) to improve the release profile of the previous implant. (Fig. 12) Their study demonstrated that, the sustained release of ganciclovir could be successfully modulated by blending high and low molecular weight poly(D,L‑lactide). The ganciclovir release duration could be controlled from several months to one year, depending on the blending ratio of poly(D,L‑lactide). The modified implants composed of poly(D,L‑lactide)-70000 and poly(D,L‑lactide) ‑5000 with a ratio of 80/20 could maintain the ganciclovir concentration in the vitreous of the pigmented rabbit eye within the range of ED₅₀ without a significant burst in the late release phase. Their results suggest that the blended implants are useful for the intraocular controlled drug delivery, over a period of several months to one year to treat cytomegalovirus retinitis. [13]

Figure 12

III.3 MAGNETICALLY CONTROLLED RELEASE

The above studies involve using materials to deliver drugs at constant or decreasing release rates. However, increases in release rate might be useful in diseases like diabetes, for example. To examine this, Langer et al. developed magnetic polymeric composites. Their thought was to use an elastic material such as ethylenevinyl acetate copolymer as the polymer matrix and add magnetic beads and powdered drug to it. Without an external magnetic field, drug is released by slow diffusion through pores. However, when an oscillating magnetic field is activated, these pores are compressed, and more drugs are released [8]. Eventually, they suggested that, by designing a triggering system in the form of a special wristwatch‑like device that could be programmed or connected to a biosensor, release could be activated on demand.

To test their idea, an oscillating magnetic field was designed from two Plexiglas disks placed vertically on top of each other and separated by several inches. A motor was attached to the bottom disk. The top disk was stationary, and the bottom one, which contained magnets, rotated. When exposed to the oscillating magnetic field, the polymer‑magnet composites displayed release rates up to 30 times higher than those without the magnetic field. Release rate could be controlled by magnetic field strength and frequency. [8]

To test this process in vivo, they designed a polymer system with a 2‑year supply of insulin and a small magnet. These systems were implanted subcutaneously in diabetic rats, and the same rotating device was used. The rats were placed in small cages on the stationary top disk, while the bottom disk rotated, giving the rats 20‑min exposures to the oscillating magnetic field. Blood sugars were lowered to a near‑normal level, when the magnetic field was applied.

More recently, they designed a microchip to deliver drugs (Figure 13). The microchip is made of silicon and contains multiple (up to 1000) drug reservoirs, each covered with a thin gold film. The chip is made using photolithography, chemical vapor deposition, and reactive ion etching; the reservoirs are filled via inkjet printing or microinjection. By applying approximately 1 V selectively to any individual reservoir, the gold was dissolved electrochemically, thereby releasing entrapped drug. Proof of principle release studies; have been conducted with single or multiple model drugs. Langer et al. lastly suggested that this microchip can potentially be regulated by remote control, or eventually a biosensor might be attached to create a smart delivery system. [8]

Figure 13

IV. FURTHER SUGGESTIONS

The limitations of traditional modes of drug administration, namely non-specific, non-steady drug concentration levels, are forcing scientists to search on new modalities. A large number of scientific group, therefore, are studying on controlled release of drugs for the treatment of several diseases. These search subjects vary from controlled release devices for root canal therapy [14], to controlled protein release from polymers [15], or usage of chemical compounds other than polymers, such as rotaxanes [16]. But, amongst them, studies on controlled release cancer therapy is the most important one, in my opinion, since traditional cancer treatments do not give satisfactory results but just elongate patient life in an undesirable quality because of side effects, pain and so lowered physicology.

Designing materials, that can be used for the controlled release of a class of drug would be more applicable, but to design a controlled release system for a specific drug to treat a specific disease sounds to be more succesful although it would have the disadvantage of higher cost. For such an improved specific design, first the polymer that is going to degrade optimally should be synthesized. The mechanism of polymer degradation and degradation time should be taken into account.

Targetting drugs to diseased organs or tissues by using magnetic field is one of the approachs developed in last years [17]. If the magnetic field can also be used for controlled release, one example of magnetically controlled release was explained by Langer et al. [8], this would be a very effective cancer treating system. Of course, designing such a system requires a large knowledge and a strong background on the polymer chemistry, medicine, and biomedical engineering.

V. REFERENCES

[1] Kydonieus; A. F., “Controlled Release Technologies: Methods, Theory, and Applications”, volume I, CRC Press, 1980, 2-6.

[2] Dash; A. K., Cudworth; G. C., “Therapeutic Applications of Implantable Drug Delivery Systems”, Journal of Pharmacological and Toxicological Methods, 1998, 40, 1-12, and the references therein.

[3] Babincova; M., Sourivong; P., Chorvat; D., Babinec; P., “Laser Triggered Drug Release From Magnetoliposomes”; Journal of Magnetism and Magnetic Materials, 1999, 194, 163-166.

[4] Hirayama; F., Uekama; K., “Cylodextrin-based Controlled Drug Release System”; Advanced Drug Delivery Reviews, 1999, 36, 125-141.

[5] Fung; K. F., Saltzman; W. M., “Polymeric Implants for Cancer Chemotherapy”; Advanced Drug Delivery Reviews, 1997, 26, 209-230, and the references therein.

[6] Chasin; M., Langer; R., “Biodegradable Polymers as Drug Delivery Systems”, Marcel Dekker Inc., 1990, 261.

[7] Dhanikula; A. B., Panchagnula; R., “Localized Paclitaxel Delivery”; International Journal of Pharmaceutics, 1999, 183, 85-100, and the references therein.

[8] Langer; R., “Biomaterials in Drug Delivery and Tissue Engineering: One Laboratory’s Experience”, Accounts of Chemical Research, 2000, 33, 94-101.

[9] Jain; R. A., “The Manufacturing Techniques of Various Drug Loaded Biodegradable Poly(lactide-co-glycolide) Devices”, Biomaterials, 2000, 21, 2475-2490 and the references therein.

[10] Risbud; M. V., Hardikar; A. A., Bhat; S. V., Bhonde; R. R., “ pH-Sensitive Freeze-Dried Chitosan-Polyvinyl Pyrrolidone Hydrogels as Controlled Release System for Antibiotic Delivery”, Journal of Controlled Release, 2000, 68, 23-30.

[11] Han; J. H., Krochta; J. M., Kurth; M. J., Hsieh; Y., “Lactitol-Based Poly(ether polylol) Hydrogels for Controlled Release Chemical and Drug Delivery Systems”, Journal of Agricultural and Food Chemistry, 2000, 48, 5278-5282.

[12] Kuijpers; A. J., Engbers; G. H. M., Meyvis; T. K. L., de Smedt; S. S. C., Demeester; J., Krijgsveld J., Zaat; S. A. J., Dankert; J., Feijen; J., “Combined Gelatin‑Chondroitin Sulfate Hydrogels for Controlled Release of Cationic Antibacterial Proteins”, Macromolecules, 2000, 33, 3705-3713.

[13] Kunou; N., Ogura; Y., Yasukawa; T., Kimura; H., Mlyamotob; H., Hondab; Y., Ikada; Y., “Long‑term sustained release of gancielovir from biodegradable scleral implant for the treatment of cytomegalovirus retinitis”, Journal of Controlled Release, 2000, 68, 263-271.

[14] Huang; J., Wong; H. L., Zhou; Y., Wu; X. Y., Grad; H., Komorowski; R., Friedman; S., “In vitro Studies and Modelling of a Controlled- Release Device for Root Canal Therapy”, Journal of Controlled Release, 2000, 67, 293-307.

[15] Tbata; Y., Ikada; Y., “ Protein Release from Gelatin Matrices”, Advanced Drug Delivery Reviews”, 1998, 31, 287-301.

[16] Chiu; S. H., Rowan; S. J., Cantrill; S. J., Glink; P. T., Garrell; R. L., Stoddart; J. F., “A Rotaxane Like Complex with Controlled Release Characteristics”, Organic Letters, 2000, 2, 3631-3634.

[17] Viroonchatapan; E., Sato; H., Ueno; M., Adachi; I., Tazawa; K., Horikoshi; I., “Magnetic Targetting of Thermosensitive Magnetoliposomes to Mouse Kivers in an in situ on-Line Perfusion System”, Life Sciences, 1996, 58, 2251-2261.

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