Biomedical Engineering Reference
In-Depth Information
Surfactants
LbL self-assembly
FIGURE 10.7
Scheme of direct coating of polyelectrolytes onto small molecule drug crystals. Surfac-
tants could be used to impart charges onto noncharged drug crystals and followed by LbL self-assembly of
polyelectrolytes.
charged amphipathic polymers could be used fi rst (Figure 10.7). It has been demonstrated that a
model drug pyrene crystals exposed to didodecyldimethylammonium bromide (positively charged)
showed a zeta-potential of
50 mV, while sodium dodecyl sulfate (negatively charged) dispersed
another model drug fl uorescein diacetate crystal exhibited a value of
+
50 mV [157]. Once charged,
polyanions and polycations can be further applied for shell encapsulation (Figure 10.7).
The next important step is to choose appropriate polymers for LbL self-assembly. Natural poly-
mers are preferable because of their excellent biocompatibility and biodegradability. Commonly
used hydrophilic materials in pharmaceutical applications include polysaccharides (e.g., CH), poly-
peptides (e.g., poly[aspartic acid]), and proteins (e.g., gelatin). Besides, synthetic biocompatible and
biodegradable polymers are also good candidates. For example, poly(α-[4-aminobutyl]-l-glycolic
acid) (PAGA), a novel poly-l-lysine analog, containing both a degradable ester linkage and a posi-
tively charged backbone, provides good biocompatibility and degradability for gene and interleukin
delivery [158,159]. Poly(β-amino ester), a cationic and degradable polymer, when built with heparin
multilayers, showed tunable drug release upon hydrolytic degradation of the polymers [160].
The drug release rate can be directly controlled by varying the number of layers of coating
[153,154,161,162]. In a model drug study experiment, fl uorescein dye microcrystals were used as
templates, and multilayers of PSS and PAH were self-assembled on particle surface [116]. For 8 to
18 layers, the permeability value decreased from 7
-
10
-
9
m/s. The permeability can be
converted into a diffusion coeffi cient (
D
) by means of multiplying the permeability with the shell
wall thickness. Assuming 3 nm for each individual polyelectrolyte layer, the calculated diffusion
coeffi cients ranged from 1.7
×
10
-
9
to 2
×
10
-16
m
2
/s. The Mohwald group has discovered that there
is a limitation when increasing fi lm thickness to reduce the shell permeability, that is, permeability
can no longer be changed after a certain number of layers have been assembled [162]. For example,
after 16 bilayers of PAH/PSS assembly on ibuprofen drug crystals were achieved, further shell
thickness increase did not reduce shell permeability. This phenomenon can be explained by fi lm
homogenization after coating certain number of layers [162]. Basically, pores or defects may exist in
the fi lms at the beginning of LbL self-assembly because of incomplete coverage by polyelectrolytes.
It can be expected that the permeation through these pores or defects will give rise to the permeabil-
ity of the fi lms. With the number of coating layers increased, the pores or defects could be reduced
or completely closed. As a result, the fi lm permeability decreases and eventually reaches a constant
value when the fi lms are structurally homogeneous.
Another parameter to control shell permeability is choosing different materials for shell assem-
bly. On encapsulation with 15 bilayers of polysaccharides, CH and dextran sulfate, the half-release
time of ibuprofen (
t
1/2
) was about twice as bare drug microcrystals at pH 7.4 [154]. Only fi ve bilayers
of PAH/PSS coating prolonged the half-release time about 6 times longer than bare microcrystals
[162]. Using gelatin led to even much slower drug release; it has been demonstrated that at pH 7.4,
six bilayers of gelatin and PSS prolonged half drug release time about 300 times compared with
uncoated furosemide microcrystals [153]. Similarly, slower drug release from gelatin-containing
×
10
-16
to 1
×
