Biomedical Engineering Reference
In-Depth Information
from patient to patient. On the other hand, the levels of
enzymatic activity may vary widely not only from pa-
tient to patient but also between different tissue sites in
the same patient. Thus polymers that undergo hydro-
lytic cleavage tend to have more predictable in vivo
erosion rates than polymers whose degradation is me-
diated predominantly by enzymes. The latter polymers
tend to be generally less useful as degradable medical
implants.
phenoxy)propane as the monomeric starting material.
Likewise, devices made of PGA erode faster than iden-
tical devices made of the more hydrophobic PLA,
although the ester bonds have about the same chemical
reactivity toward water in both polymers.
The observed bioerosion rate is further influenced by
the morphology of the polymer. Polymers can be classi-
fied as either semicrystalline or amorphous. At body
temperature (37 C) amorphous polymers with T g above
37 C will be in a glassy state, and polymers with a T g
below 37 C will in a rubbery state. In this discussion it is
therefore necessary to consider three distinct morpho-
logical
Factors that influence the rate
of bioerosion
states:
semicrystalline,
amorphous-glassy,
and
amorphous-rubbery.
In the crystalline state, the polymer chains are densely
packed and organized into crystalline domains that resist
the penetration of water. Consequently, backbone hy-
drolysis tends to occur in the amorphous regions of
a semicrystalline polymer and at the surface of the
crystalline regions. This phenomenon is of particular
importance to the erosion of devices made of poly
( L -lactic acid) and PGA which tend to have high degrees
of crystallinity around 50%.
Another good illustration of the influence of the poly-
mer morphology on the rate of bioerosion is provided by
a comparison of poly( L -lactic acid) (PLLA) and poly( D , L -
lactic acid): Although these two polymers have chemically
identical backbone structures and an identical degree of
hydrophobicity, devices made of PLLA tend to degrade
much more slowly than identical devices made of poly( D ,
L -lactic acid). The slower rate of bioerosion of PLLA is due
to the fact that this stereoregular polymer is semi-
crystalline, while the racemic poly( D , L -lactic acid) is an
amorphous polymer.
Likewise, a polymer in its glassy state is less permeable
to water than the same polymer when it is in its rubbery
state. This observation could be of importance in cases
where an amorphous polymer has a glass transition
temperature that is not far above body temperature
(37 C). In this situation, water sorption into the polymer
could lower its T g below 37 C, resulting in abrupt
changes in the bioerosion rate.
The manufacturing process may also have a significant
effect on the erosion profile. For example, Mathiowitz
and co-workers (Mathiowitz et al, 1990) showed that
polyanhydride microspheres produced by melt encas-
pulation were very dense and eroded slowly, whereas
when the same polymers were formed into microspheres
by solvent evaporation, the microspheres were very
porous
Although the solubilization of intact polymer as well as
several distinct mechanisms of chemical degradation have
been recognized as possible causes for the observed
bioerosion of a solid, polymeric implant, virtually all cur-
rently available implant materials erode because of the
hydrolytic cleavage of the polymer backbone (mechanism
III in Fig. 3.2.7-2 ). We therefore limit the following dis-
cussion to solid devices that bioerode because of the hy-
drolytic cleavage of the polymer backbone.
In this case, the main factors that determine the
overall rate of the erosion process are the chemical sta-
bility of the hydrolytically susceptible groups in the
polymer backbone, the hydrophilic/hydrophobic char-
acter of the repeat units, the morphology of the polymer,
the initial molecular weight and molecular weight dis-
tribution of the polymer, the device fabrication process
used to prepare the device, the presence of catalysts,
additives, or plasticizers, and the geometry (specifically
the surface area to volume ratio) of the implanted device.
The susceptibility of the polymer backbone toward
hydrolytic cleavage is probably the most fundamental
parameter. Generally speaking, anhydrides tend to hy-
drolyze faster than ester bonds that in turn hydrolyze
faster than amide bonds. Thus, polyanhydrides will tend
to degrade faster than polyesters that in turn will have
a higher tendency to bioerode than polyamides. Based on
the known susceptibility of the polymer backbone
structure toward hydrolysis, it is possible to make pre-
dictions about the bioerosion of a given polymer.
However, the actual erosion rate of a solid polymer
cannot be predicted on the basis of the polymer back-
bone structure alone. The observed erosion rate is
strongly dependent on the ability of water molecules to
penetrate into the polymeric matrix. The hydrophilic
versus hydrophobic character of the polymer, which is
a function of the structure of the monomeric starting
materials, can therefore have an overwhelming influence
on the observed bioerosion rate. For instance, the erosion
rate of polyanhydrides can be slowed by about three
orders of magnitude when the less hydrophobic sebacic
acid is replaced by the more hydrophobic bis(carboxy
(and
therefore
more
water
permeable)
and
eroded more rapidly.
The preceding examples illustrate an important
technological principle in the design of bioeroding de-
vices: The bioerosion rate of a given polymer is not an
unchangeable property, but depends to a very large
Search WWH ::




Custom Search