Implants are very small pellets composed of drug

substance only without excipients.

They are normally about 2-3 mm in diameter and

are prepared in an aseptic manner to be sterile.

Implants are inserted into a superficial plane

beneath the skin of the upper arm by surgical

procedures, where they are very slowly absorbed

over a period of time.

Implant pellets are used for the administration of

hormones such as testosterone.

IMPLANTABLE CONTROLLED DRUG DELIVERY SYSTEMS

The capsules may be removed by surgical

procedures at the end of the treatment period.

Biocompatibility need to be investigated, such as

the formation of a fibrous capsule around the

implant and, in the case of erosion-based devices

there is the possible toxicity or immunogenicity of

the byproducts of polymer degradation.

The Implantable controlled drug delivery system

achieved with two major challenges.

1) by sustained zero-order release of a therapeutic

agent over a prolonged period of time.

This goal has been met by a wide range of techniques,

including:

Osmotically driven pumpsOsmotically driven pumps

Matrices with controllable swellingMatrices with controllable swelling

diffusiondiffusion or erosion ratesor erosion rates

2) By the controlled delivery of drugs in a pulsatile or

activation fashion.

These systems alter their rate of drug delivery in

response to stimuli including the presence or absence

of a specific molecule, magnetic fields, ultrasound,

electric fields, temperature, light, and mechanical

forces.

Such systems are suitable for release of

therapeutics in non-constant plasma concentrations

as in diabetes.

This goal has been met by two different

methodologies:

A delivery system that releases the drug at a

predetermined time or in pulses of a predetermined

sequence.

A system that can respond to changes in the local

environment.

IMPLANTABLE CONTROLLED DRUG DELIVERY SYSTEMS IN A PULSATILE FASHION

Theoretical pulsatile release from a triggered-system.

IMPLANTABLE CONTROLLED DRUG DELIVERY SYSTEMS

IN A SUSTAINED ZERO-ORDER CONTINUOUS RELEASE

In membrane permeation-type controlled drug

delivery, the drug is encapsulated within a

compartment that is enclosed by a rate-limiting

polymeric membrane.

The drug reservoir may contain either drug

particles or a dispersion of solid drug in a liquid or a

solid type dispersing medium.

The polymeric membrane may be made-up from a

homogeneous or a heterogeneous nonporous

polymeric material or a microporous or

semipermeable membrane.

The drug release by diffusion (dQ/dt) from this type

of implantable therapeutic systems should be constant

and defined by:

dt

dQ=

dm

R

PP

C11

Where:

CR is the drug concentration in the reservoir

compartment and Pm are the permeability coefficients of

the rate-controlling membrane

Pd the permeability coefficients of the diffusion layer

existing on the surface of the membrane, respectively.

Pm and Pd depend on the partition coefficients for the

interfacial partitioning of drug molecules from the

reservoir to the membrane and from the membrane to

the aqueous diffusion layer, respectively.

Example Levonorgestrel ImplantsThese are a set of six flexible, closed capsules

of a dimethylsiloxane/methylvinylsiloxane copolymer,

each containing 36 mg of the progestin

levonorgestrel.

They are inserted through a 2 mm incision in the

mid-portion of the upper arm in a fan-like pattern.

This system provides long-term (up to 5 years)

reversible contraception.

Diffusion of the levonorgestrel through the wall of

each capsule provides a continuous low dose of

progestin.

A technique that depend on sequential release of

drugs which fabricated as polymer matrix with

multilayer alternating drug-containing and spacer

layers.

Pre-programmed Delivery Systems

The polymer matrix is commonly surrounding

impermeable shell, which permitting release of the

entrapped drug only after degradation of this polymer

matrix.

For degradation of this polymer matrix to occur, the

polymer matrix must be susceptible to hydrolysis or

biodegradation by a component in the surrounding

media.

)A) Schematic of a multilayered pulsatile delivery system with one face exposed to the local environment.

(B) Schematic of a cylindrical multilayered delivery system with two open faces.

I.System that controlling drug

release by environmental pH

Using polyanhydrides as the spacer layers and the

drug containing layer as poly[(ethyl glycinate)

(benzly amino acethydroxamate)phosphazene]

(PEBP)

The polyanhydrides and PEBP layers were

compression molded to form a multilayered

cylindrical core, which was then coated with a

poly(lactide-co-1,3-trimethylene carbonate) film over

all surfaces except for one face of the device.

The hydrolysis of PEBP is highly dependent on the

pH of the surrounding media, dissolving much more

rapidly (1.5 days) under neutral and basic conditions

(pH 7.4) but in acidic conditions (pH 5.0) digradad

over 20 days.

The degradation products of polyanhydrides create

an acidic environment within the delivery device,

preventing the rapid hydrolysis of the PEBP and

result in slow drug release until all of the

polyanhydride layer has been eroded.

Using hydrogels that have differing susceptibilities

to enzymatic degradation.

Pulsatile release can be achieved with a model

system that uses the enzymatic degradation of

dextran by dextranase to release insulin in a

controlled manner.

A delivery vehicle can be fabricated by covering

poly(ethylene glycol)-grafted (embedded) dextran

(PEG-g-Dex) and unmodified dextran layers in a

silicone tube.

II. System that controlling drug release by environmental enzymes

The drug is loaded into the PEG-g-Dex layers while

dextran is material for the spacer layer.

The introduction of PEG into a dextran solution

containing a drug causes the formation of a two-

phase polymer when the dextran is cross-linked.

The drug is partitioned into the PEG phase,

resulting in drug release that is erosion-limited

instead of diffusion-limited.

Closed-loop delivery systems

Closed-loop delivery systems are those that are

self-regulating.

They are similar to the programmed delivery

devices in that they do not depend on an external

signal to initiate drug delivery.

However, they are not restricted to releasing their

contents at predetermined times. Instead, they

respond to changes in the local environment, such as

the presence or absence of a specific molecule.

Glucose-Sensitive Systems

Several strategies are used for glucose-responsive

drug delivery.

1. pH Dependent systems for glucose-stimulated drug

delivery

2. Competitive binding

1. pH Dependent systems for glucose-stimulated

drug delivery

As insulin is more soluble under acidic conditions,

Incorporating glucose oxidase into a pH-responsive

polymeric hydrogel enclosing insulin solution will

result in a decrease in the pH of the environment

immediately surrounding the polymeric hydrogel in

the presence of glucose as a result of the enzymatic

conversion of glucose to gluconic acid.

)A) Diagram of a glucose-sensitive dual-membrane system.

(B) The membrane bordering the release media responds to increased

glucose levels by increasing the permeability of the membrane

bordering the insulin reservoir.

A copolymer of ethylene vinyl acetate (EVAc)

containing g glucose oxidase immobilized on cross-

linked poly- acrylamide. and insulin solution . the

insulin release rate will be altered in response to

changes in the local glucose concentration.

The release rate of insulin returned to a baseline

level when the glucose was remove.

A dual-membrane system

sensing membrane is placed in

contact with the release media,

while a PH barrier membrane is

positioned between the sensing

membrane and the insulin reservoir.

As glucose diffuses into the hydrogel , glucose

oxidase catalyzes its transport to gluconic acid, thereby

lowering the pH in the microenvironment of the PH

membrane and causing swelling .

Gluconic acid is formed by the interaction of glucose

and glucose oxidase, causing the tertiary amine groups

in the PH- membrane to protonated and induce a

swelling response in the membrane.

Insulin in the reservoir is able to diffuse across the

swollen barrier membrane.

Decreasing the glucose concentration allows the pH of

barrier membrane to increase, returning it to a more

collapsed and impermeable state .

2. Competitive binding methodology depending on the fact that

concanavalin A (Con A) a glucose-binding lectin, can

bind both glycosylated insulin and glucose.

Glycosylated insulin (G-insulin) bound to Con A can

be displaced by glucose, thus release the drug from

system.

In this systems immobilized Con A -Glycosylated

insulin encapsulated with a polymer (sepharose beads

) , release only occurs at sufficiently high glucose

concentration .

as Con A immobilized has a lower binding affinity

for glucose than for G-insulin, preventing release at

low glucose levels.

Hydrogels formed by mixing Con A and (G-insulin)

with copolymers as acrylamide .

hydrogel will undergo a reversible gel–sol phase

transition in the presence of free glucose due to

competitive binding between the free glucose and

Con A.

G-insulin acts as a cross-linker for the Con A chains

due to the presence of four glucose-binding sites on

the molecule, but competitive binding with glucose

disrupts these cross-links, making the material more

permeable and thus increasing the rate of drug

delivery.

Sol–gel phase transition in polymers crosslinked with Con A.

Similar systems have been developed that use the

interaction between an antibody and an antigen to

control the release of a drug in the presence or

absence of the antigen.

A hydrogel held together by the interaction of

polymer-bound antigen to polymer-bound antibody

will swell in the presence of free antigen due to the

competitive binding of bound antibody to free

antigen, reducing the number of crosslinks in the

hydrogel and thus increasing the rate of drug

delivery in proportion to the antigen concentration.

Open-loop Delivery Systems Open-loop delivery systems are not self-regulating,

but require externally generated environmental

changes to initiate drug delivery.

These can include magnetic fields, ultrasound,

electric fields, temperature, light, and mechanical

forces.

Open-loop delivery systems may be coupled to

biosensors to obtain systems that automatically

initiate drug release in response to the measured

physiological demand.

1. Magnetic Field One of the first methodologies to achieve an

externally controlled drug delivery system is the use

of an magnetic field to adjust the rates of drug

delivery from a polymer matrix.

A magnetic steel beads embedded in an EVAc

copolymer matrix that is loaded with the drug.

An oscillating magnetic field ranging from 0.5 to

1000 gauss cause increased rates of drug release.

The rate of release could be altered by changing the

amplitude and frequency of the magnetic field.

The increased release rate was caused by mechanical

deformation due to magnetic movement within the

matrix.

During exposure to the magnetic field, the beads

oscillate (swing) within the matrix, creating compressive

and tensile forces which acts as a pump to (squeezing)

push an increased amount of the drug molecule out of

the matrix.

2. Ultrasound Ultrasound stimulus can be used to adjust drug

delivery by directing the waves at a polymer or

hydrogel matrix.

Where drug release can be increased 27-fold from

an EVAc matrix during exposure to ultrasound.

Increasing the strength of the ultrasound resulted

in a increase in the amount of drug released (1 W/cm

for 30 min).

The principle depends on that sound cavitation

occurred by ultrasonic irradiation at a polymer–liquid

interface forms high-velocity jets of liquid directed at

the polymer surface that are strong enough to

release away material at the surface of the polymer

device, causing an increase in the erosion rate of the

polymer .

Also the sound cavitation enhances mass transport

at a liquid–surface interface.

Electric Field

Electric current signal can be used to activate drug

delivery.

The presence of an electric current can change the

local pH which initiate the erosion of pH-sensitive

polymer and the release of the drug contained in

polymer matrix.

Polymers as poly(methacrylic acid) or poly(acrylic

acid) can be dissolved at pH>5.4

A 5 mA electric current resulted in drug delivery due

to the production of hydroxyl ions at the cathode,

which raised the local pH, disrupting the hydrogen

bonding between the comonomers.

In the absence of the electric stimulus, drug release

was negligible.

Humans can tolerate direct current densities of

under 0.5 mA/cm for up to 10 min; therefore no

visible skin damage was observed.

Temperature

Thermally-responsive hydrogels and membranes can

be used for pulsatile delivery of drugs.

Temperature sensitive hydrogels have a lower critical

solution temperature (LCST), a temperature at which a

hydrogel polymer undergo a phase change. In which

transition of extended coil to the uncross-linked

polymer an can be occurred .

This phase change is based on interactions between

the polymer and the water surrounding the polymer.

Thermally sensitive hydrogel systems can exhibit

both negative controlled release, in which drug

delivery is stoped at temperatures above the LCST,

and positive controlled drug delivery, in which the

release rate of a drug increases at temperatures

above the LCST.

N-Isopropylacrylamide (NIPAAm) is a commonly

used thermosensitive polymer with an LCST of 32 °C.

Thermally sensitive materials exhibiting negative

thermally controlled drug delivery.

When the temperature of the hydrogel is held below its

LCST, the most thermodynamically stable configuration

for the free (non-bulk) water molecules is to remain

clustered around the hydrophobic polymer. When the

temperature is increased over the LCST, the collapse of

the hydrogel is initiated by the movement of the

clustered water from around the polymer into bulk

solution. Once the water molecules are removed from

the polymer, it collapses on itself in order to reduce the

exposure of the hydrophobic domains to the bulk water.

Thermally sensitive materials exhibiting positive

thermally controlled drug delivery.

A copolymer of NIPAAm and acrylamide (AAm) is an

example of such a material. The hydrophilic AAm

increases the LCST of the copolymer as well as

reducing the thickness and density of the outer layer

formed when the temperature of the hydrogel is

raised above its LCST.

Upon collapse, the hydrogel will push out soluble

drug held within the polymer matrix

5. Light

The interaction between light and a material can be

used to adjust drug delivery.

This can be accomplished by combining a material

that absorbs light at a desired wavelength and a

material that uses energy from the absorbed light to

adjust drug delivery.

Near-infrared light has been used to adapt the

release of drugs from a composite material fabricated

from gold nanoparticles and poly(NIPAAm-co-AAm)

When exposed to near-infrared light, the nanoshells

absorb the light and convert it to heat, raising the

temperature of the composite hydrogel above its

LCST (40 °C(. This in turn initiates the

thermoresponsive collapse of the hydrogel, resulting

in an increased rate of release of soluble drug held

within the polymer matrix.

6. Mechanical force

Drug delivery can also be initiated by the

mechanical stimulation of an implant.

Alginate hydrogels can release included drugs in

response to compressive forces of varying strain

amplitudes.

Free drug that is held within the polymer matrix is

released during compression; once the strain is

removed the hydrogel returns to its initial volume.

This concept is similar to squeezing the drug out of

a sponge.