Mechanism of Drug Release From Implantable Devices

By: Pharma Tips | Views: 3546 | Date: 04-Jul-2011

Drug release from most implantable devices is controlled by any one of six different mechanisms discussed below. Although an attempt is made here to cover the most importable types it is not possible to cover all the mechanisms under investigation.

Drug release from most implantable devices is controlled by any one of six different mechanisms discussed below. Although an attempt is made here to cover the most importable types it is not possible to cover all the mechanisms under investigation.

Diffusion controlled
These devices are based on Flick’s law of diffusion which states the rate of transfer of a diffusing substance through unit area of a section. In this case the rate of release is controlled by diffusion of drug through a polymeric membrane. In general nonerodible diffusion-controlled drug delivery system work best for drugs with molecular weight of 1000 Daltons or less. Domb et al have shown that the essential parameters affecting permeability of peptides through a hydrogel are the volume of solute and the water content of the membrane which are correlated to its pore size. Diffusion-controlled devices can be further classified into membrane-permeation controlled matrix-controlled and micro reservoir-dissolution controlled.

Membrane-permeation controlled
In membrane-permeation controlled devices the drug reservoir is surrounded by a membrane (coating) and because of the presence of he two distinct drug-reservoir and membrane phases these are known as heterogeneous devices when the device containing a highly hydrophilic drug is placed in the aqueous dissolution medium water penetrates the coating and dissolves the drug and the concentrated drug solution diffuses out through the polymeric membrane. The release rate of the drug is controlled by the diffusion rate of drug solution through the polymeric membrane. The rate of drug release dM/dt through a spherical membrane-permeation controlled system with saturated reservoir is given by equation.

DM/dt    = 4DK(C1 – C2) ab / b-a………….(1)

Where, D is the diffusivity of drug unit thickness of polymer k is the partition coefficient (ratio of solubility of drug in the polymer divided by the solubility of drug in the surrounding medium) of drug across the polymer membrane c2 is the concentration of drug inside the sphere c2 is the concentration of drug in the surroundings a is the inner radius of the coat and b is the outer radius of the coat.
Mathematical models describing the effect of device geometry on the release pattern have been developed. An important advantage of the reservoir system as shown by eq. (1), is that at steady state zero-order drug release is possible. It the drug elimination rate is constant during therapy, a constant drug-plasma level may be achieved. On the other hand any disruption or crack in the membrane could lead to the release of a large amount of drug (dose dumping) with possible toxic effects in the patient. Furthermore reservoir systems are generally more expensive to manufacture.

Diffusion-controlled reservoir systems can also be based on a biodegradable polymer. In this case, the drug is encapsulated in a biodegradable polymer and the release rate is determined by the principles governing membrane-permeation controlled systems. Only after the entire drug is exhausted from the device does the polymer undergo significant erosion and eventually dissolves.
Currently marketed and approved membrane reservoir systems for implantation therapy include the Ocusert an ocular insert for controlled pilocarpine release for the treatment of glaucoma; Progestasert. a T-shaped contraceptive intrauterine device the controlled progesterone release and Norplant which is comprised of six small injectable cylinders for controlled release of levonorgestrone for contraception.


Matrix controlled

In matrix-controlled devices the drug is uniformly distributed (dissolved of dispersed) throughout the polymer and hence these are known as homogeneous devices. In the presence of dissolution medium drug at the surface dissolves first and is released in the dissolution medium. In many cases the dissolved drug creates a depletion boundary separating the empty or drug-depleted polymer from the drug-loaded polymer matrix. Water penetrates the channels and a pore created by drug depletion and dissolves the drug
at the depletion boundary. The drug release rate is controlled by the diffusivity barrier provided by the empty matrix which increases in thickness with time. This increased thickness results in a decrease in drug release rate with time. For a matrix system that is exposed to the dissolution medium on all the sides the surface area of the inward-moving depletion boundary decrease resulting in a decrease I in drug release rate which depends on the device-geometry.

Mathematical equation describing release have been reported which predict that in general the drug release rate is expected to decrease with time. In case of damage to the device though the drug release rate may increase slightly significant dose dumping is not expected. Therefore these devices have a safety design superior to that of the membrane-controlled system. Furthermore matrix systems are less expensive to manufacture. It has been shown that manipulation of the shape or drug distribution may allow a constant delivery rate.

Microreservoir dissolution-controlled
In these devices the drug reservoir is made of a suspension of solid drug particles in an aqueous solution of a water miscible polymer forming millions of microscopic drug reservoirs in a polymer matrix. the device is coated with a rate-controlling membrane to further modify the drug release rate. Among the other factors the release rate is dependent on the solubility of drug in the liquid compartment and on the polymer matrix. Mathematical relationships for the control of drug release have been described.

The Syncro-Mate-M implant provides an example of this type of device. It is a cylindrical implant fabricated by dispersing the drug reservoir (a suspension of norgestomer in an aqueous solution of PEG 400) in silicone elastomers. This emulsion is placed into medical-grade silicone tubing and polymerized in situ. The tubing is then cut to obtain a cylinder-shaped device with open ends. This device is implanted subcutaneously in he earflap of livestock for about 20 days delivery of norgestomet for control and synchronization of estrus and ovulation. Other examples include the Nitrodics system and a dual-release vaginal contraceptive ring system.

Chemically controlled
Chemically controlled drug delivery systems regulate the drug release rate by a chemical reaction with the polymer. Their principal advantage is that in contrast to a nonbioerodible system the polymer is dissolved and absorbed by the body and there is no need for surgical removal of the device at the completion of drug delivery. However the fate of these polymeric products in the body must be carefully observed and rigorous testing is required to confirm the safety of the polymer. The two predominant mechanisms for chemically controlling drug release are bioerosion and pendent chain.

Bioerosion
The bioerosion or biodegradation systems (both terms are used interchangeably in this article) involve breakdown of the polymer into small water-soluble molecules. Bioerosion-controlled devices are matrix controlled with uniform drug distribution inside the polymer. As the polymer is broken down water comes in contact with the drug leading to its dissolution and release. Depending on the hydrophilicity of the polymer and device properties such as porosity and the presence of water-soluble components water may penetrate throughout the device or come in contact only with the surface. In the former case polymer erosion starts throughout the matrix; these devices are known as bulk eroding. Although the drug release initially occurs predominantly from the erosion of the polymer from the surface eventually the entire matrix may break down and most of the remaining drug could be released in a burst. On he other hand if the polymer is hydrophobic and water does not penetrate inside the device, erosion occurs only on the surface; these devices can be called surface eroding. The drug release rate from a surface eroding polymeric matrix with uniform drug distribution is given by eq.

dM/dt = KS………………..(2)

Where, K is a constant related to the drug concentration in the matrix and the rate of polymer erosion and S is the surface area of the system. In general, due to erosion of polymer, the surface area of devices of the most common shapes decreases over time, rate of this decrease is geometry dependent. For systems with a high initial surface are to-volume ratio such as a slab the surface area and hence the release rate, decreases only slightly. On the other hand with devices of spherical geometry where the surface area-to-volume ratio is the lowest the surface area and hence release rate decreases rapidly with time. Therefore, matrix drug-release devices with shapes of slab geometry may provide less deviation in drug release over time than those with spherical geometry.

In practice because of imperfections such as pin holes, the presence of water-soluble drugs and polymer hydrophilicity, bioerodible delivery systems are most commonly bulk-erosion controlled. The drug release from these systems depends on a combination of diffusion bulk erosion and surface erosion, making a theoretical prediction of drug release rate much more difficult. An important advantage of this type of system is that the drug-release profile may be controlled by manipulation of the size and shape of the device amount of drug loading addition of other excipients and the intrinsic degradation rate and molecular weight of the polymer.

Pendent chain
The other mechanism for chemically controlled release of drug is known as the pendent-chain system where the drug is attached to the polymer backbone by a labile chemical linkage. In the presence of water or enzymes the labile linkage breads to release the drug. The pendent chain may be water soluble or insoluble; a water-soluble backbone may serve as a drug carrier to a specific cell or organ where the drug is released by metabolism. Insoluble pendent chains on the other hand serve as a depot from which the drug is slowly released. In either case after completion of the drug release the polymer-drug linkage. Varying the hydrophilicity of the polymer backbone and the device geometry can control the rate of linkage degradation and therefore the drug release. An important disadvantage of this system compared to bioerosion-controlled systems is that
since the drug is covalently linked to the polymer the drug-polymer conjugate may be viewed as a new chemical entity by the regulatory agencies and extensive safety testing may be needed.

Solvent activated

Solvent-activated systems release active agents because of controlled penetration of a solvent into the device; they may be controlled by swelling or osmotic pressure.

Swelling controlled
Swelling-controlled systems are similar to matrix-type devices except that the dispersed drug is immobilized inside a glassy polymer and therefore there is no diffusion of drug. When this device is placed in water the outer polymer region begins to swell, resulting in relaxation of the polymer chains. This allows the otherwise locked drug to diffuse outward. Therefore two fronts are observed: one moving inward, separating the polymer in the glassy stare from the rubbery state and the second moving outward separating the swollen rubbery polymer from the surrounding aqueous medium. The drug release is determined by the rate of relaxation of the chains that unlock the drug.

Osmotically controlled
In one example of an osmotically controlled system an osmotically active agent such as water-soluble salt is placed inside a rigid semi permeable polymer housing, which is separated from the drug compartment by a movable partition. The semi permeable housing draws water inside by osmosis, leading to an increase in volume and exertion of pressure on the movable partition. The partition, in turn pushes the drug out of the compartment through a delivery orifice or cannula. Thus, the drug delivery rate is controlled by the mass movement of water across the semi permeable membrane.
 
Osmotic pump provide a predictable, zero-order release rate independent of the physicochemical properties of the drug and therefore afford an excellent means of evaluating effects of long-term, zero-order administration into animals. Alza corp. markets several implantable osmotic pump devices. Alzet osmotic pumps have been extensively evaluated for local as well as systemic drug delivery in may animal models. Delivery rates from these pumps range from 1 to 10 L/h, and a release duration from three days to four weeks. The smallest of these osmotic pump, model 10003d, weighs about 350 mg and is designed for small research animals, weighing about 20 g. although these pumps are very useful in providing constant drug release in animals, they have found minimal therapeutic application in humans because of the need for surgical implantation and removal lack of external regulation and in situ refillability. However a variation of the osmotic pumps where a drug or and osmotically active agent is surrounded by a semi permeable membrane containing a single laser-drilled hole (the OROS system is being used for oral delivery of drugs such as phenylpropanolamine (Acutrim, Ciba0 and nifedipine (pericardia-XL, Pfizer)

Externally regulated
These systems have the important advantage that the drug-delivery rate cat be externally increased on demand even after the device has been implanted. Four predominant techniques have been evaluated with externally modulated implant: magnetically controlled ultrasonically activated thermally activated and electrically controlled.

Magnetically controlled
In magnetically controlled drug-delivery systems the drug and magnetic beads are uniformly dispersed inside semi elastic polymer matrix made of a nonbiodegradable polymer such as ethylene-vinyl acetate copolymer (EV Ac). When the device is placed in a dissolution medium the drug release follows matrix diffusion control. However, when the device is placed in a magnetic field, the magnetic beads attempt to a align with the applied magnetic field including a torque on the magnet and a slight rearrangement of the polymer. In an oscillating magnetic field, the beads tend to oscillate compressing and expanding the polymer in the process. This is proposed to result in a pulsatile flow of the
dissolution medium through the pores in the elastic polymer and along the concentration gradient of the drug resulting in an increase in the drug-release rate. This effect of the oscillating magnetic field is further enhanced by an increase in polymer elasticity (e.g., by increasing the vinyl acetate content) and the frequency and strength of the magnetic field.
 
Ultrasonically activated
In these systems the drug is uniformly distributed inside a polymer and an external ultrasonic field is applied to activate drug release. They have been evaluated for both nonbiodegradable polymers (EVAc) and biodegradable polymers [polyesters, polyanhydrides, polyglycolides, polylactides and sebacic acid]. In the case of biodegradable polymers application of ultrasound increased the drug release as well as the polymer degradation rate. It is believed that the cavitation induced by the ultrasonic waves may be partially responsible for this effect as reduced polymer degradation and drug release was observed in a degassed buffer. In both the biodegradable and nonbiodegradable polymer systems the drug release rate was controlled by the intensity frequency and duration of the ultrasound.
For in vivo evaluation devices made of EVAc and insulin or polyanhydrides and p-amminohippurate were implanted in rats and an ultrasonic applicator head was placed over the treated area for the delivery of ultrasonic waves to the implant. The rate of drug delivery was increased with no histologically detectable damage to the rat skin.

Thermally activated
A series of thermosensitive hydrogels that show significant swelling changes in water in response to temperature have been prepared and evaluated. These polymers respond to temperature change based on the Flory-Huggins theory that a change in temperature affects hydrogen bonding which, in turn, affects swelling. A linear correlation is observed between the diffusion coefficient for entrapped drug and polymer swelling. Based on their origin by thermosinsitive interaction these polymers can be classified into those base on polymer-water interactions and those based on polymer-polymer interactions.

In the first group, a series of poly (N-alkyl substituted acrylamides) have been evaluated for the effect of cross-linking density on the temperature dependence of swilling. At low temperatures the polymer is in an unswollen state due to the increased hydrogen bonding interaction. This swilling leads to increased solute diffusion and hence drug release. Thus thermosensitivity of the polymer network has been shown to primarily due to polymer-water interactions and hydrophobic interactions of the side groups.

In the second group two interacting polymers with repulsive or attractive polymer-polymer interaction are combined to form a hydrogel. With an increase in temperature, swilling increases as the polymer responds not only to the polymer-water interactions but also to increased polymer-polymer interactions.

Hoffman demonstrated the applicability of hydrogels based on N-isopropylacrylamide (NIPAAm) or N-isopropylacrylamide-methacrylic copolymers with methylenebis (acrylamide) (MBAAm) in providing temperature-responsive release of vitamin B12. By manipulation of temperature, pulsatile release of Indomethacin and insulin has been reported.

Electrically controlled
Electrically controlled systems provide drug release by the action of an applied electric field on a rate-limiting barrier membrane or a solute thus modulating its transport across it. Grimshaw reported four different mechanisms for the transport of proteins and neutral solutes across

hydrogel membranes:

Electrically and chemically induced swelling of a membrane to alter the effective pore size and permeability
Electrophoretic augmentation of solute flux within a membrane
Electroosmotic augmentation of solute flux within the membrane and
Electrostatic partitioning of charged solutes into charged membranes.

Propanolol hydrochloride delivery devices containing a drug reservoir with a pair of electrodes placed across a poly (2-hydroxyethyl methacrylate) (PHEMA) membrane, cross-linked with ethylene glycol (1 % v/v) have been evaluated for modulated drug delivery. Controlled and predictable propranolol hydrochloride release in response to the electric field was observed with a linear relationship between current and propranolol hydrochloride permeability. Variables for this mechanism have been extensively evaluated for Transdermal delivery (iontophoresis) systems.

Self-regulated
These are biofeedback-controlled system, where the drug release rate is dependent on the body’s need for the drug at a given time. From a therapeutic viewpoint these systems may come closest to duplicating the release from a gland such as the pancreas. A variety of mechanisms have been employed to obtain self-regulated delivery.

Ionic strength and pH responsive
These devices take advantage of the fact that polymers containing weakly acidic or basic side groups develop a charge in alkaline or acidic pH respectively. In a cross-linked water-insoluble polymer, this results in water uptake and corresponding swelling of the polymeric membrane with opening of molecular pores and increased drug release rate. Siegel demonstrated the application of cross-linked polymeric gels (methyl methacrylate-N, N-dimethylaminoethl methacrylate co polymer) in drug delivery. IN this case, the polymer is unionized and hydrophobic in neutral pH. A reduction in the pH of the gel leads to ionization of the gel-forming polymer with subsequent swelling which results in molecular pores for release of drug.Thus,this system has shown to effectively control caffeine release at different pH.In terms of practical usefulness.This concept has been extended to glucose sensitive release of insulin as describe below.

Glucose responsive
Immobilized Glucose Oxidase. Glucose Oxidase catalyses a reaction between glucose and oxygen in the body fluids to form gluconic acid, which reduces the pH of the microenvironment. This is related to the concentration of gluconic acid and hence glucose. The insulin-release systems based on glucose Oxidase utilize this drop in pH to trigger an increased release.

Heller reported on application of Ph-sensitive biodegradable polyorthoesters in which insulin is immobilized. The biodegradable polymer matrix is coated with a hydrogen containing immobilized glucose Oxidase. Glucose diffusing into the hydrogel is oxidized to gluconic acid causing a drop in pH. This increases the erosion rate of the polymer releasing entrapped insulin. Thus, insulin release in modulated by the concentration of glucose in the surrounding fluids.

Langer took advantage of the fact that like many other insulin has a pH of minimum solubility around its isoelectric point. They used modified insulin (with three additional lysine molecules) change the isoelectric point of insulin from around 5.0 to 7.4 thus at physiological pH insulin has the lowest solubility. In the presence of glucose and immobilized glucose Oxidase the resultant drop in pH leads to an increase in insulin solubility with an increase in release rate.
Another such system employs drug encapsulated in special phospholipids vesicles that change structure and permeability with changes in ph temperature or glucose concentration.

Competitive binding
Brownlee and Cerami discovered the principle of competitive binding. It involves the preparation of glycosylated insulin which combines with Concavalin A (Con A), a

Carbohydrate- binding protein. Con A is immobilized on sepharose beads and the glycosylated insulin is bound to it. The beads are encapsulated in a suitable membrane which is permeable to both glucose and insulin. Blood glucose diffuses into the device and replaces insulin from the con a-insulin complex by competitive binding. The free insulin diffuse out of the rate-controlling membrane and is thus released in an active and more stable against aggregation than commercial insulin. This device has been shown to produce glucose-dependent insulin release both in vitro and in vivo. However, the release of insulin has been found to be nonlinearly dependent on glucose.
The device was implanted intraperitoneally in dogs whose pancreas had been removed and blood-glucose levels were compared with those of normal and diabetic dogs after administration of 500-mg/kg dextrose bolus dose. The implanted dogs responded to blood-glucose levels in a manner similar to that of normal dogs. The levels were maintained in the dog for the two days of in vivo evaluation.

Urea responsive
Heller was first to demonstrate the feasibility of a self-regulated hydrocortisone delivery system responding to the presence of urea. The device of disks containing hydrocortisone incorporated into a biodegradable polymer (n-hexyl half-easer of methyl vinyl ether and maleic anhydride) with pH dependent degradation. This disk is coated with a hydrogen containing immobilized urease. In physiological-buffer base line hydrocortisone release is obtained by the hydrolysis of the polymer and diffusion of drug. In the presence of urea the enzyme urease increases the ph of the microenvironment by converting urea into ammonium bicarbonate and ammonium hydroxide. This increase in pH results in increased hydrolysis of the biodegradable polymer and increased hydrocortisone release. The latter was shown to be proportional to the concentration of urea present.
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