Biomedical Engineering Reference
In-Depth Information
The bonding mechanism of dense HA implants ap-
pears to be very different from that described above for
bioactive glasses. The bonding process for HA implants is
described by Jarcho (1981). A cellular bone matrix from
differentiated osteoblasts appears at the surface, pro-
ducing a narrow amorphous electron-dense band only
3-5
m
m wide. Between this area and the cells, collagen
bundles are seen. Bone mineral crystals have been iden-
tified in this amorphous area. As the site matures, the
bonding zone shrinks to a depth of only 0.05-0.2
m
m
(
Fig. 3.2.10-2
). The result is normal bone attached
through a thin epitaxial bonding layer to the bulk implant.
TEM image analysis of dense HA bone interfaces show an
almost perfect epitaxial alignment of some of the growing
bone crystallites with the apatite crystals in the implant.
A consequence of this ultra-thin bonding zone is a very
high gradient in elastic modulus at the bonding interface
between HA and bone. This is one of the major differ-
ences between the bioactive apatites and the bioactive
glasses and glass-ceramics. The implications of this dif-
ference for the implant interfacial response to Wolff's law
is discussed in Hench and Ethridge (1982, Chap. 14).
rapidly resorbed than HA. TCP has four polymorphs,
a
bg
and super a
.
The
g
polymorph is a high pressure
phase and the super a polymorph is observed at tem-
peratures above approximately 1500
C. Therefore the
most frequently observed polymorphs in bioceramics
are a and b-TCP. X-ray diffraction studies indicate that
the b polymorph transforms to the a polymorph at
temperatures between 1120
C and 1290
C (Gibson
et al.
, 1996).
In the 1980s, the idea of a new bone susbtitute ma-
terial was introduced and the materials were referred to
as calcium phosphate bone cements. These materials
offer the potential for
in situ
molding and injectability.
There are a variety of different combinations of calcium
compounds (e.g.,
a
-TCP and dicalcium phosphate) that
are used in the formulation of these bone cements, but
the end product is normally based on a calcium-deficient
HA (Fernandez
et al.,
1998,1999a, b).
All calcium phosphate ceramics biodegrade to varying
degrees in the following order: increasing rate HA.
The rate of biodegradation increases as:
1.
Surface area increases (powders
>
porous solid
>
dense solid)
2.
Crystallinity decreases
3.
Crystal perfection decreases
4.
Crystal and grain size decrease
5.
There are ionic substitutions of CO
2
3
,Mg
2
þ
, and
Sr
2
þ
in HA
Factors that tend to decrease the rate of biodegradation
include (1) F
substitution in HA, (2) Mg
2
þ
substitution
in b-TCP, and (3) lower b-TCP/HA ratios in biphasic
calcium phosphates.
Resorbable calcium phosphates
Resorption
or
biodegradation
of
calcium
phosphate
ceramics is caused by three factors:
1.
Physiochemical dissolution, which depends on the
solubility product of the material and local pH of its
environment. New surface phases may be formed,
such as amorphous calcium phosphate, dicalcium
phosphate dihydrate, octacalcium phosphate, and
anionic-substituted HA.
2.
Physical disintegration into small particles as a result
of preferential chemical attack of grain boundaries.
3.
Biological factors, such as phagocytosis, which causes
a decrease in local pH concentrations (de Groot and
Le Geros, 1988).
Ideally, one would wish for a replacement material to
be slowly resorbed by the body once its task of acting as
a scaffold for new bone has been completed. Degrada-
tion or resorption of calcium phosphates
in vivo
occurs
by a combination of phagocytosis of particles and the
production of acids. However, when selecting a resorb-
able material for implantation, care must be taken to
match the rate of resorption with that of the expected
bone tissue regeneration. Where the solubility of cal-
cium phosphates is higher than the rate of tissue re-
generation, they will not be of use to fill bone defects.
As mentioned previously, the rate of dissolution in-
creases with decreasing calcium-to-phosphorus ratio,
and consequently, TCP, with a Ca:P ratio of 1.5, is more
Clinical applications of HA
Calcium phosphate-based bioceramics have been used in
medicine and dentistry for nearly 20 years (Hulbert
et al.
, 1987; de Groot, 1983, 1988; de Groot
et al.
,
1990; Jarcho, 1981; Le Geros, 1988; Le Geros and Le
Geros, 1993). Applications include dental implants,
periodontal treatment, alveolar ridge augmentation, or-
thopedics,
maxillofacial
surgery,
and
otolaryngology
(
Table 3.2.10-6
).
Most authors agree that HA is bioactive, and it is
generally agreed that the material is osseoconductive,
where osseoconduction is the ability of a material to
encourage bone growth along its surface when placed in
the vicinity of viable bone or differentiated bone-forming
cells. A good recent review of
in vitro
and
in vivo
data for
calcium phosphates has been prepared by Hing
et al.
(1998), who observed that there are a large number of
''experimental parameters,'' including specimen, host,
