Controlled release of biologically active compounds from bioerodible polymers

Controlledrelease of biolo$ically active compoundsfrom bioerodible polymers cJ. Heller Polymer

Sciences Department,

Received

8 June 1979; revised 1 October

SRI lnterna tional, Menlo Park, CA 94025,

USA

1979

This article reviews the controlled release of biologically active agents by the erosion or chemical degradation of a polymer matrix into which the agent is incorporated. Chemically bound active agents and work on steroid release from cholesterol implants are not covered. The mechanisms of polymer erosion discussed are: cross-linked scission; hydrolysis, ionization or protonation of pendant groupts; backbone cleavage. Drug release studiesare dealt with under each of these headings.

It is now generally recognized that the controlled release of biologically active agents to a local environment can be achieved by means of one of three general methodologies: (1) diffusion through a rate~ontrolling membrane, (2) use of osmotically regulated flow, and (3) release controlled by the erosion or chemical degradation of a matrix into which the active agent is incorporated’. Each of these methodologies offers certain unique characteristics which determine the design of specific therapeutic systems. Thus, methodology (1) allows construction of drug delivery devices that release therapeutic agents by zero order kinetics and where rate of delivery can be readily adjusted by changing the rate-limiting membrane and/or membrane thickness and area. Methodology (2) allows construction of devices that not only release their contents by zero order kinetics but are also able to sustain high delivery rates not normally available with membranemoderated devices. Meth~ology (3) allows construction of drug delivery devices that have a predetermined life span and need not be removed from the site of action once their drug delivery role has been completed. Drug release from bioerodible polymers finds use in both topical applications and systemic applications. Both uses demand that the polymer degrade to nontoxic products, and polymers used in systemic applications must also degrade to low-molecular-weight fragments that can be readily eliminated or metabolized by the body. In topical applications, retainment of high molecular weight of the degradation products is desirable, since in this way no unnecessary systemic absorption of the polymer will occur, and toxicological hazards are thus reduced. The purpose of this article is to present a comprehensive review of methodology (3) where active agents are released to a surrounding aqueous environment by solubilization of the polymer matrix induced by the aqueous environment. The review is limited to devices in which the active agent is dissolved or dispersed in a polymer, and does not 0142-9612/80/010051-07

$02.00

@

1980

IPC

Business

cover the important work in which the active agent is chemically bound to the polymer and is released to the surrounding medium by hydrolysis of a bond between the active agent and the polymer chain2s3; nor does it cover the extensive work of Kincl and coworkers on the release of steroids from cholesterol implants4. For this review, it is convenient to systematize polymer erosion according to the three mechanisms shown in Figure 7, where @ denotes a hydrolytically unstable bonds. In general terms, Mechanism I encompasses watersoluble polymers that have been insolubilized by hydrolytically unstable crosslinks; Mechanism II includes polymers that are initially water-insoluble and are solubilized by Mechanism I

--*hcmim

--

A

II

,

A

0

Mechanem

I

C

IE

---_--Figure

1

Schematic

representation

of degradation

mechanisms;

@ denotes a hydrolytica/& unstable bond represents a hydrophobjc substituent and A B + C represents hydrolysis, ionization or protonation

Press Biomaterials

1980,

Vol t January

51

3ioarod~bfe

polymers:

J. Heller

hydrolysis, ionization, or protonation of a pendant group; and Mechanism III includes hydrophobic polymers that are converted to small water-soluble molecules by backbone cleavage. Clearly, these three mechanisms represent extreme cases, and erosion by a combination of mechanisms is possible.

MECHANISM

HO

q-~-N”-CH,-NH--q 2

Time(days) Figure 2 Release of hydrocortisone gela tin matrix

1980,

0 2-+E-NHZ+

0 H-h-l

cleavage

In these systems, water-soluble polymers are insolubilized by means of hydrolyti~ally unstable crosslinks. Consequently, the resulting matrix is highly hydrophilic and completely permeated by water. Since the active agent is in an aqueous environment, its water solubility becomes an important consideration, and compounds with appreciable water solubility will be rapidly leached out, independent of the matrix erosion rate. There are two general applications in which erodible hydrogels are useful in the controlled delivery of active agents. In the first the active agent has extremely low water solubility, and in the second the active agent is a macromolecule that is entangled in the hydrogel matrix and cannot escape until a sufficient number of crosslinks have cleaved and matrix crosslink density has been reduced. The first application is illustrated in figure 2, which shows the release of a highly water-insoluble drug, hydrocortisone acetate, from a gelatin matrix crosslinked with formaldehyde5. As indicated by the first-order dependence, release is by diffusion with little contribution by matrix erosion, Because the drug is very water insoluble, useful release over many days is achieved. Such a device could be used when zero-order kinetics are not important and removal of the expended device is not convenient or desirableb. Illustrative of the second application are many examples in which water-soluble macromolecules have been immobilized in hydrogels by physical entanglement7. However, the intent of most of these studies was to achieve long-term immobilization of enzymes or antigens; the slow diffusional escape and/or slow hydrolysis of the matrix with consequent liberation of the entrapped macromolecules was generally regarded as undesirable. Two studies, however,

Biomaterials

cl

0

I

Solubilization by cr~iink

52

recognized the utility of erodible hydrogeis for providing sustained delivery of entrapped macromolecules. Both studies took advantage of the hydrolytic instability of crosslinks formed in vinyl polymers by using N, N’-methylenebisacrylamide as a comonomer. Hydrolysis of the crosslinks proceeds as follows:

Vo/ 1 January

acetate

from a crosslinked

where -R- represents the vinyl polymer chain. In the first study8, bovine pancreatic insulin was immobilized in a hydrogel prepared from acrylamide and 2% N, N’-methylenebisacrylamide. Slow release of insulin from the hydrogel was inferred because insulin-containing hydrogel implants sustained diabetic animals for at least a few weeks. It is not clear from the study how much insulin was released by diffusion and how much by cleavage of crosslinks. In the second study9, cu-chymotrypsin was immobilized in a hybrogel prepared from N-vinyl pyrrolidone and N, No-methylenebisacrylamide. It was found that, by varying the N, N’-methylenebisacrylamide concentration from 0.1 to 1 .O w/w % with respect to N-v~nylpyrrol idone, hydrogels with dissolution times of several days to practically insoluble could be prepared. However, release of cu-chymotrypsin did not correlate well with hydrogel dissolution time, presumably because of diffusional escape. To prevent diffusional escape, a-chymotrypsin was acylated with acryloyl chloride and then chemically incorporated in the hydrogel by copolymerization with N-vinyl pyrrolidone and N, N’-methyfenebisacrylamide. Release of cY-chymotrypsin from the resulting hydrogels was then found to correlate more closely with hydrogel dissolution times.

MECHANISM

II

Solubilization by hydrolysis, ionization or protonation of pendant groups Systems in this category include all polymers that are initially water-insoluble but become water-soluble as a consequence of hydrolysis, ionization, or protonation of pendant groups. Because no backbone cleavage takes place, the solubilization does not result in any significant changes in molecular weight. The major emphasis in the development of these materials has been on enteric coatings. These are coatings designed to be insoluble in a certain pH environment, usually the stomach, and then to dissolve abruptly in an environment of a different pH, such as the intestines. Usually these polymers are applied to pills as protective coatings and do not produce steady, sustained release of therapeutic agents. However, by using mixtures of enteric coatings, each with a different disintegration time, it is possible to prolong the action of therapeutic agents’*. Literature on enteric coatings is much too voluminous for detailed review, but the coatings can be grouped into three categories according to their dissolution mechanism: (1) dissolution by side group hydrolysis, (2) dissolution by ionization of a carboxylic acid function and (3) dissolution by protonation of amine functions.

Bioerodible

Dissolution category

by side group hydrolysis.

are represented

and maleic

Materials

by copolymers

in this

of vinyl

progressively

monomers

anhydride:

polymer,

more

ionization

resulting

same argument grouping

is necessary

in increasingly

holds

degrees

the degree of esterification, polymer HO

:

-CH,mCH-FH-FH-A A 0

Recently

: XH;CH-?H-YH-

A 0 0

copolymers

it has been shownt7

of methyl

vinyl

in a constant

pH environment,

H”

H”

incorporated

therein

of hydrocortisone films

X is OR or H. In the anhydride

insoluble, because

form,

but on exposure affect

time

before

initiation

before

dissolution

substituent

increases

R in the vinyl

increases,

and decreases

increases.

enteric

ether

generally

water-insoluble, functions

become

acidic

acetate

media

phthalate

based on cellulose

Currently

used

and can be represented form

they

of the carboxylic coatings

They

esterified

anhydride

in aqueous succinate

and maleic anhydride

It was found

that

acid

in aqueous

they

are insoluble.

0.25 pH units, atoms

they

and changes

coatings

polymer

chain

into

the polymer; the alkyl

hence

group

With

of ionization

sharp,

which

about

of carbon

readily

by considering

to drag the

relatively

that

cortisone.

correlated

with from

provides result

pH is low.

amount

to solubilize As the size of

so does the hydrophobicity,

and

film

in water

at a pH

took

solution

place

for hydro-

of size of alkyl

group.

polymer

can be directly and was, in fact,

and strong

release rates

dependence

on

Since in all experiments

with that

total drug

on

between

eroded

release. All drug

linearity

group

esters measured

correlation

polymer

drug

dissolution,

release and polymer

of pH on rate of polymer

release of hydrocortisone

partial

for

diffusional

concomitantly.

The effect hence

evidence

was independently

of drug released

also coincided

The

release

over several days.

drug

again it can be assumed occur

strong

of the matrix

release, distance

measuring

dissolved

and drug

rate for a series of partial

again show excellent

erosion

erosion

the aqueous

None was found

and drug

of polymer

and for negligible

no dissolution analysing

erosion

erosion

ester is shown

dispersed

in Figure

dependence

of erosion

the eroding

medium

erosion

and

in the n-butyl

5. The date show

a clear

rate and drug release on the pH of

and, as expected,

a progressive

decrease

in rate as the critical dissolution pH is approached. The partial ester copolymer also has been used recently that

small ester

is sufficient

loo

016 0

of

exhibit

and below

necessary

the dissolution

increases,

ether

of the copolymers.

solution.

a low degree

vinyl

copolymers

the number

carboxyls

polymer

The latter

of drug

was determined

loss of the device.

a drug-containing

the size of the alkyl

characteristically

are soluble

can be understood

of ionized

of both

by placing

and periodically

have also been investigated14-16.

with

in the ester side group

groups,

of methyl

This pH range is quite

This behaviour the number

verified

polymer

the amount

weight

mechanism

derived

on the

bases. Enteric

esterified

these materials

a pH range above which

the total

of the device

polymer

have also been

copolymers or partially

in which

a surface-erosion

depletion

Partially ethylene

linearity

kinetics.

Each pair of points

and the amount

from

can,

rate and the rate

the wash solution

over the lifetime

are

described13. and maleic

by U.V. measurements was calculated

pH. esterified

anhydride

half-ester

drug.

Figure 4 shows the effect

are based on

are insoluble

of the free acid groups

acetate

into

excellent

and maleic zero-order

at pH 7.4. Because of the linear

used enteric

dissolve

device

low enough

increases. groups.

a separate

partially

dissolution

the dispersed

by the device

the

the dissolution

release hydrocortisone

by excellent

release of the drug.

environalso increases

water-soluble.

because

radical,

In general,

of the copolymer

in unionized

phthalatet2.

but,

and lag

dissolution

this type,

While

The most widely cellulose

portion

but on ionization

they

A number

process”.

of carboxyl

represent

as polyacids.

soluble

group.

dissolution

environment

by ionization

released

as the size of the alkyl

The rate of polymer

coatings

become

as the pH of the aqueous

as the pH of the aqueous Dissolution

they

of the dissolution

represents

are water-

of the anhydride

the rate of polymer

time

ment

to water

of the hydrolysis

of variables

the polymers

The higher

hydrophobic

release for n-butyl

containing

the

pH. The

the same ester

that

ether

cob

Figure 3 shows polymer

where

having

the higher

co8

to solubilize

of esterification.

the more

and consequently

J. Heller

high dissolution

for polymers

but different

polymers:

as a model

for a bioerodible

releases a therapeutic

of a specific cortisone methyl

external

molecule18.

was incorporated vinyl

drug delivery

agent in response

ether-maleic

into

In this model, an n-hexyl

anhydride

system

to the presence hydro-

half ester of a

copolymer

and the

100

1

IO

20

1

30 Time (h)

I

40

I

I

50

60

Figure 3 Rate of polymer dissolution and rate of release of hydrocortisone for the n-butyl half-ester or methyl vinyl ether-maleic anhydride copolymer containing 10 wt-% drug dispersion. lo), drug release; (nl, polymer dissolution. Reproduced from J. Appl. Polym. Sci. 1978, 22, 1991 by permission of John Wiley and Sons Inc., New York

“0.

IO

20

30

40

50

60

70

80

90

Time (run) Figure 4 Effect of size of ester group in halfesters of methyl vinyl ether-maleic anhydride copolymers on rate of erosion at pH 7.4. Reproduced from J. Appl. Polym. Sci. 1978, 22, 1991 by permission of John Wiley and Sons inc., New York

Biomaterials

1980,

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53

Bioerodible

25

polymers:

J. Heller

A

.

IO Oo

IO

20

30

40 50 60 70 60 Time (min) Figure 5 Effect of erosion medium pH on erosion rate of half-asters of methyl vinyl ether-meleic anhydride copolymers. A = pH 5.5, B=pH5.75, C=pH6.0, D=pH6.5, E=pH7. Reproducedfrom J. Appl. Polym. Sci. 1978, 22, 1991 with permission of the copyright owner

polymer-drug mixture fabricated into discs. These discs were then coated with a hydrogel containing immobilized urease. In a medium of constant pH and in the absence of external urea, the rate of release of hydrocortisone was that normally expected for that polymer at the given pH. In the presence of external urea, ammonium bicarbonate and ammonium hydroxide were generated within the hydrogel, which accelerated polymer erosion and hence drug release. Drug delivery rate increase was proportional to the amount of external urea and was reversible; that is, when external urea was removed, the rate of drug release gradually returned to its original value. The effect of urea on polymer erosion and drug release is shown in Figure 6. Date presented thus far show that the partial esterpolymer system is a useful polymer matrix for zero-order drug delivery; it shows the desired surface erodibility, and the erosion rate at any given pH can be adjusted by selecting the appropriate ester alkyl group. The usefulness of this polymer system for delivery of a therapeutic agent in vivo has been investigated by fabricating hydrocortisone-containing films in small circles and placing them in the lower forniceal cul de sac of New Zealand 100

0

1

20

40

60

00

100 120 140 Time (h)

160 160

200 220

240

Figure 6 Rate of hydrocortisone release at 35°C from a nhexyl half-ester of a copolymer of methyl vinyl ether and maleic anhydride at pH 6.25 in he absence and presence of external urea. A = polymer + hydrogel, IO-lrn urea; B =polymer + hydrogel, lo-2m urea; C = polymer + hydrogel, no urea. Reproduced from J. Pharm. Sci. 1979, 68,919 with permission of the copyright owner

54

Biomaterials

1980,

Vol 1 January

Figure 7 rabbits

20

30

40 Time (h)

In vivo release of hydrocortisone

50

60

70

1

from ocular inserts in

rabbit@. Kinetic plots were obtained by placing weighed devices in eyes and removing the devices at desired time intervals. The amount of drug released was determined by measuring residual drug in the devices and subtracting that from the known original amount. The results expressed in terms of polymer erosion are shown in Figure 7. Clearly, the devices are highly functional and release hydrocortisone by excellent zero-order kinetics. Furthermore, considering that each point represents a different device and a different rabbit eye, there is very little scatter, indicating a high degree of reproducibility. Polymer dissolution by protonation of amine functions. Materials in this group are, in effect, reverse enteric coatings; they are insoluble in water and alkali but soluble in acids. Although no drug-release studies have been described, suggested uses have been as moisture-resistant medicament coatings that will release their contents in the stomachlg,) and as veterinary medicament coatings to allow passage through the stomach of ruminants and subsequent release in the abomasum20. A material typical of this group is cellulose acetate N, N-diethylaminoacetate, prepared by the amination of cellulose acetate chloroacetate with diethylamine21. Another material was prepared by the addition of amines to crotonic acid esters of cellulose22.

MECHANISM

III

Solubilization

by backbone cleavage

This category includes all water-insoluble polymers that undergo hydrolytic backbone cleavage and are solubilized by conversion to small, water-soluble molecules. A major driving force for the development of these materials was a search for absorbable sutures superior to catgut. From these studies have evolved two synthetic materials that were suitable for surgical implants: poly (lactic acid)23,24 and poly(glycolic acid)25. The first demonstration of the utility of poly(lactic acid) as a bioerodible implant capable of sustained release of a therapeutic agent was described about 10 years ago26, and since then many publications have dealt with the sustained release of pharmaceuticals from bioerodible implants. For this review, the literature will be discussed according to the type of pharmaceutical agent delivered: (I) delivery of narcotic antagonists, (2) delivery of contraceptive steroids, and (3) delivery of others. Delivery of narcotic antagonists. The release of cyclazocine from poly(lactic acid)27 and the release of Naltrexone

8ioerodjbie

polymers:

.I. Heller

Time (days) Figure 9 In vitro release of 3H-Naitrexone base from uncoated rods of 75125 poly IL f+j-iacticcoglycolic acid) into 37°C. buffered. pH7 solution as a functian of drug loading. x, 80% w/w n. base; W, 70% w/wn. base; A, 60% w/wn. base; l,50% w/wn. base. Reproduced with permission from Life Sci 1975, 17. 1877 @ 1975 Pergamon Press Ltd.

to

20

30 40 50 60 Trne (days) Figure 8 Cumulative amounts of cyclazocine excreted from composites of poly(lactic acid) implanted or injected in rats. Amounts have been corrected to account for unrecovered drug, lo), (Ol, and (A), film; (01, small particles; ( O), film with drug sealed into poly (lactic acid) envelope. Reproduced with permission Chem. 1973, 16,897 @American Chemical Society

from .I. Med.

from poly(lactic acid)28 have been described in two studies. In the first study, composites of poly(lactic acid) and labelled cyclazocine were prepared and implanted in rats as 4 cm* films, as ground films having particle sizes failing within No. 25/35 sieves, and as 6.2 cm* poly(lactic acid}-cyclazocine composites enveloped in poly(lactic acid] containing no drugs. In vivo release was followed by monitoring the radioactivity of excreted urine. The results, shown in Figure 8 were not those expected; cyclazocine is released by diffusion, and therefore release from small particles with their greater surface area should be much faster than release from films. It was postulated27 that the in vjvo results are influenced by varying degrees of inflammation and oedema. This may be reasonable, because in vitro release data show the expected much faster release from particles than from films. Since the polymer was said to bioerode in about 62 days and after 55 days, only 30% of the drug was released, a large burst in drug delivery should take place shortly thereafter. In the second study2*, tritiated Naltrexone was incorporated into poly{lactic acid), and the composite was ground and then injected into rats. Drug release, as measured by urinary excretion, levels off at about 25% after about 30 days. Again, there was a considerable disparity between in vivo and in vitro release, this time attributed to tritium exchange in the body. Release of cyclazocine microencapsulated in poly (lactic acid) has also been described.29 However, because of macroscopic defects in the capsule wall, all cyclazocine was released in vivo between 14 and 17 days. The permeability of cyclazocine through solvent-cast poly(lactic acid) was measured to yield a value of 2.9 and 3.0 x 10-t’ cm*/sec. Using an average of these values, it was calculated that a defect-free capsule should release 50% of its contents in about 28.5 months.

The in vitro release of Naltrexone and Naltrexone pamoate from poly(lactic acid) microcapsules has also been described3*. Again, the microcapsules contained defects, and about 50% of the contents were released after the first few hours of extraction. Copolymers of lactic and glycolic acid have also been used in a bioerodible delivery system for Naltrexone and Naltrexone pamoate3’. The in vitro release rates of Naltrexone from rods of 75/25 poly(lactic/glycolic acid) copolymer as functions of drug loading are shown in Figure 9. As expected, increase in drug loading results in increased delivery rates. ln vivo release of labelled Naltrexone from 90/10 poly(lactic/glycolic acid) beads is shown in Figure 10. The release was measured as urinary excretion; after about 90 days about 50% of the drug was accounted for, The remaining drug was assumed to have been excreted in the farces. Other investigators3’ found that Naltrexone is excreted in approximately equal amounts in the urine and in the farces. Delivery of contraceptive steroids. d-Norgestrel was incorporated into poly(lactic acid)~and its release rate followed both in vitro and in vivo33. It was found that films of the polymer containing 33% Norgestrel released the steriod at a relatively constant rate of 3 kg/day per cm* for over 80 days. When similar films were implanted subcutaneously in rats, the initial release was about 5.5 &g/day per cm*; by 80 days release had declined to about 3 @g/day

j

1~~ 10

, , , , , , ,i 20

30

40

50

60

70

80

90

Time (days past lrn~~~tion) Figure 70 Cumulative percent of 3H, implanted as 3H-Naltrexone base (33% w/w) in 9O~lOpoiy lLi+l-iacticcoqlycolic acid) beads, recovered in the urine of mice. Reproduced with permission from Life Sci 1975, 17, 1877 @ 1975 Pergamon Press Ltd.

Biomaterials

1980,

Vol 1 January

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Bioerodible

polymers:

J. Heller

Time (days1

Time (doysf Figure 1 I Cumulative percent norethisterone released in vitro from particles containing 20% by weight nore thisterone. iOj90f8Ou; uncoated (0) 180-256~; uncoated (01 180-250~; coated fusing a 10% poly_DL-lactic acidhenzene solution). From Contraception 1976, 13,375. Reproduced with permission of the copyright owner

per cm*. Biodegradation of the matrix was slower than drug release. In another study3 14C-labelled 20 wt% Norethisterone was incorporated into poly (L(+)-lactic acid) and the composite was cryogenically ground and separated into 90- to 180-p particles and 180- to 250~fi particles. The l80- to 250-p particles were overcoated with the benzenesoluble po!y(DL -lactic acid). The in vitro release of the steriod from all three particle-size formulations is shown in Figure ?I. Release from the larger particles was faster than that from the smaller particles, which is contrary to what might be expected because rate of diffusional release is directly proportional to surface area. Release from coated samples is considerably lower than that from uncoated samples, This is as expected; the steriod must diffuse through a rate-controlling membrane. Studies of in viva release, as measured by urinary excretion, show an apparent zero-order delivery with an initial burst and another burst around day 90, which is attributed to dissolution of the polymer matrix. Comparison of actual in vim release kinetics for all three samples is difficult because the steriod is excreted in both urine and faeces, and in the uncoated samples only the urinary steroid was measured. Furthermore, either because of cumulative error or errors in original specific activity, mass balance is poor. However, the results indicate fairly linear and sustained steroid release for about three months; a typical plot is shown in Figure 12. In vivo and in vitro release of progesterone, fl-oestradiol, and dexamethasone from poly (lactic acid) beads and chips also has been described35. A considerable amount of work has been done on release of contraceptive steroids from crosslinked poly (dimethyl siloxane) implants 36. However, silicone rubber is not degradable, so the expanded device must be surgically removed. This is not desirable; consequently it is desirable to develop devices which would release steroids by membrane diffusion and the expended devices would then later bioerode. In an effort to develop biodegradable materials with permeabiiities comparable to those of silicone rubber, homo- and copolymers of e-caprolactone and DL-lactic acid were investigated as potential bioerodible membranes37. It was found that poly (e-caprolactone) and

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Biomaterials

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Figure 12 Cumulative percent 14C label from norethisterone excreted in vivo in the urine of rats from 90480 u particles of poly-L @j-lactic acid containing 20% by weight norethis~rone. From Contraception 197613.375. Reproduced with permission of the copyright owner

silicone rubber have comparable maximum steady state fluxes of progesterone at 37OC 2.2 x IO-“‘O and 0.6 x IO-” g/cm set, respectively). Poly (r&-lactic acid) had a maximum steady state flux of progesterone of 3.3 x 1o-15 g/cm sec. In another study3*, release rates of several steroids from films and capsules of homopolymers and copolymers of f-caprolactone, DL -lactic acid and glycolic acid in virro and in vivo were determined. Detailed interpretation of the data is difficult because various factors, such as simple drug diffusion, polymer erosion, morphological changes in the polymer, and changes in concentration gradient across capsule walls, all contribute to the measured erosion rate. The release of progesterone from bioerodible capsules based on glutamic acid/leucine copolymers has also been investigated3g. Devices were fabricated by first preparing progesterone-filled polypeptide rods and then coating the rods with unfilled polypeptide. Because of various fabrication difficulties, it was not possible to determine meaningful release rates. Bioerosion of the copolymers was extremely slow and complete erosion for a 50/50 copolymer was estimated to take about four years. relives of other agenrs. Although the greatest emphasis of drug delivery from bioerodible polymers has centered around narcotic antagonists and contraceptive steroids, the methodology has been applied to other drugs. A brief study described release of two anticancer drugs, cyclophosphamide and cis-dichlorodiammineplatinum, from poly(lactic acid140. Another study4’ described the release of the antimalarial drug, 2,4-diamino-6-(2naphthylsulphonyl)quinazoline (WR-158122) from a 25/75 poly(o~-lactic/ glycolic acid) copolymer in mice. Measuring radioactivity excreted into urine and faeces, sustained release through 14 weeks, was demonstrated.

DEGRADATION

MECHANISMS

The hydrolytic erosion of a solid polymer can be rationalized in terms of two extreme mechanisms. In one, referred to as homogeneous, the hydrolysis occurs at a uniform rate throughout the matrix. In the other, called heterogeneous, the process is confined to the surface of the device. Actual erosion can, of course, occur by some intermediate mechanism. In a purely homogeneous process, the matrix will remain essentially intact until all parts reach some critical

8ioerodibJ.e polymers:

degree of reaction, at which point the matrix dissolves. Drug release from such a matrix is complicated because it is a combination of diffusion and erosion. Diffusional drug release in the absence of erosion can be described by the Higuchi mode142 which proposes that the drug is initially removed from the surface regions of the polymer, and consequently a progressively thicker drugdepleted layer forms adjacent to the surface of the device. Because the remaining drug must diffuse through a progressively thickening polymer membrane, the rate of drug release declines continuously. In the presence of polymer erosion, the matrix progressively loosens up, and the permeability of the polymer to the drug increases with time. Consequently, the rate of drug release from polymers which erode either wholly or partially by bulk degradation at first shows the normal expected decline, but as the increasing polymer permeability gradually offsets this decline, the rate of drug release eventually accelerates. The heterogeneous process is a much more desirable degradation mode because it will lead to zero-order drug release, provided that diffusional release of the drug is minimal and the overall shape of the device remains essentially constant, thus maintaining constant surface area. Furthermore, such a process avoids deterioration of mechanical properties, which can take place when random chain cleavage occurs in the bulk material. In principle surface erosion can be achieved by constructing a highly hydrophobic polymer, so that water penetration and consequent bulk hydrolysis is much less probable than surface reaction. A much preferred approach is one in which a pH-sensitive reaction is selected and the interior of the matrix is buffered so that bulk hydrolysis is prevented and reaction can take place only at the surface of the device, where the buffer is neutralized. The only polymers thus far described in detail that undergo true surface erosion are the partial esters of methyl vinyl ether/maleic anhydride copolymers17‘18. However, these represent a special case because solubilization does not involve backbone cleavage, and consequently the only polymer chains than can escape into the aqueous environment are at the surface of the device. None of the described backbone-degradable polymers undergo surface erosion; rather, they undergo bulk erosion. This is not surprising, since most of them were originally designed as bioerodible suture materials and were not intended as drug carriers capable of releasing a drug by zero-order kinetics. However, work in progress here43 and elsewhere44’45 ISdevoted specifically to the development of bioerodible polymers that do undergo surface erosion and are capable of releasing incorporated drugs by zeroorder kinetics.

6

Hussain, A.A.,

7

(19761 Goldman,

8 9 10

11 12 13 14

2

3

4 5

Harris, F.W., Aulabaugh,A.E., Case, R.D., Dykes, M.K. Feld, W.A., ‘Controlled ReleasePolymeric Formulations’ (Eds. D.R. Paul and F.W. Harris) 1976, ACS Symposium

R., Goldstein,

L. and Katchalski,

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