Biomedical Engineering Reference
In-Depth Information
introduced to the market as a suture clip as well as a bone
pin marketed under the name OrthoSorb in the USA and
Ethipin in Europe.
Poly(hydroxybutyrate) (PHB), poly(hydroxyvale-
rate) (PHV), and their copolymers represent examples
of bioresorbable polyesters that are derived from mi-
croorganisms. Although this class of polymers are exam-
ples of
natural
materials (as opposed to
synthetic
materials), they are included here because they have
similar properties and similar areas of application as the
widely investigated PLA. PHB and its copolymers with up
to 30% of 3-hydroxyvaleric acid are now commercially
available under the trade name ''Biopol'' (
Miller and
Williams, 1987
). PHB and PHV are intracellular storage
polymers providing a reserve of carbon and energy to
microorganisms similar to the role of starch in plant
metabolism. The polymers can be degraded by soil bac-
teria (
Senior
et al.
, 1972
) but are relatively stable under
physiological conditions (pH 7, 37
C). Within a relatively
narrow window, the rate of degradation can be modified
slightly by varying the copolymer composition; however,
all members of this family of polymers require several
years for complete resorption
in vivo. In vivo,
PHB de-
grades to
D
-3-hydroxybutyric acid, which is a normal
constituent of human blood (
Miller and Williams, 1987
).
The low toxicity of PHB may at least in part be due to this
fact.
PHB homopolymer is very crystalline and brittle,
whereas the copolymers of PHB with hydroxyvaleric acid
are less crystalline, more flexible, and more readily pro-
cessible (Barham
et al.,
1984). The polymers have been
considered in several biomedical applications such as
controlled drug release, sutures, artificial skin, and vas-
cular grafts, as well as industrial applications such as
medical disposables. PHB is especially attractive for or-
thopedic applications because of its slow degradation
time. The polymer typically retained 80% of its original
stiffness over 500 days on
in vivo
degradation (
Knowles,
1993
).
PCL became available commercially following efforts
at Union Carbide to identify synthetic polymers that
could be degraded by microorganisms (
Huang, 1985
). It
is a semicrystalline polymer. The high solubility of PCL,
its low melting point (59-64
C), and its exceptional
ability to form blends has stimulated research on its ap-
plication as a biomaterial. PCL degrades at a slower pace
than PLA and can therefore be used in drug delivery
devices that remain active for over 1 year. The release
characteristics of PCL have been investigated in detail by
Pitt and his co-workers (
Pitt
et al.
, 1979
). The Capronor
system, a 1-year implantable contraceptive device (
Pitt,
1990
), has become commercially available in Europe and
the United States. The toxicology of PCL has been ex-
tensively studied as part of the evaluation of Capronor.
Based on a large number of tests,
3
-caprolactone and PCL
are currently regarded as nontoxic and tissue-compatible
materials. PCL is currently being researched as part of
wound dressings, and in Europe, it is already in clinical
use as a degradable staple (for wound closure).
Polyanhydrides were explored as possible sub-
stitutes for polyesters in textile applications but failed
ultimately because of their pronounced hydrolytic in-
stability. It was this property that prompted Langer and
his co-workers to explore polyanhydrides as degradable
implant materials (
Tamada and Langer, 1993
). Aliphatic
polyanhydrides degrade within days, whereas some ar-
omatic polyanhydrides degrade over several years. Thus
aliphatic-aromatic copolymers are usually employed
which show intermediate rates of degradation depending
on the monomer composition.
Polyanhydrides are among the most reactive and hy-
drolytically unstable polymers currently used as
biomaterials. The high chemical reactivity is both an
advantage and a limitation of polyanhydrides. Because of
their high rate of degradation, many polyanhydrides
degrade by surface erosion without the need to in-
corporate various catalysts or excipients into the device
formulation. On the other hand, polyanhydrides will
react with drugs containing free amino groups or other
nucleophilic functional groups, especially during high-
temperature processing (
Leong
et al.
, 1986
). The
potential reactivity of the polymer matrix toward nucle-
ophiles limits the type of drugs that can be successfully
incorporated into a polyanhydride matrix by melt
processing techniques. Along the same line of reasoning,
it has been questioned whether amine-containing bio-
molecules present in the interstitial fluid around an im-
plant could react with anhydride bonds present at the
implant surface.
A comprehensive evaluation of the toxicity of the
polyanhydrides showed that, in general, the poly-
anhydrides possess excellent
in vivo
biocompatibility
(
Attawia
et al.
, 1995
). The most immediate applications
for polyanhydrides are in the field of drug delivery. Drug-
loaded devices made of polyanhydrides can be prepared
by compression molding or microencapsulation (
Chasin
et al.
, 1990
). A wide variety of drugs and proteins in-
cluding insulin, bovine growth factors, angiogenesis in-
hibitors (e.g., heparin and cortisone), enzymes (e.g.,
alkaline phosphatase and b-galactosidase), and anes-
thetics have been incorporated into polyanhydride ma-
trices, and their
in vitro
and
in vivo
release characteristics
have been evaluated (
Park
et al.
, 1998; Chasin
et al.
,
1990
). Additionally, polyanhydrides have been in-
vestigated for use as nonviral vectors of delivering DNA
in gene therapy (
Shea and Mooney, 2001
). The first
polyanhydride-based drug delivery system to enter clin-
ical use is for the delivery of chemotherapeutic agents.
An example of this application is the delivery of bis-
chloroethylnitrosourea (BCNU) to the brain for the
