Tissue engineering - Biomedical Engineering Desk Reference - page 651

investigated for the transplantation of many cell types
including osteoblasts, chondrocytes, hepatocytes, fibro-
blasts, smooth muscle cells, and RPE. This method also
offers the possibility that genetically modified cells could
be transplanted, thereby simultaneously presenting both
the cells and the bioactive factors they produce to the site
of interest, with the potential of further enhancing re-
generation of the injured area (Blum et al., 2003). This
approach may be considered a combination of the appli-
cations of cell transplantation and delivery of bioactive
molecules, discussed later.

conditions such as temperature, pH, and heat release should
not damage implanted cells or the surrounding tissue.
The hardened material (scaffold) must be highly por-
ous and have interconnected pore structure in order to
serve as a suitable template for guiding cell growth and
differentiation. This can be achieved by combining
a porogen such as sodium chloride crystals in the injectable
paste that are eventually leached out, leaving a porous
polymer matrix. Since the leaching step occurs in vivo,
local high salt concentration may lead to high osmolarity
and tissue damage. The amount of porogen incorporated
has to be optimized to ensure biocompatibility, while
enough porosity needs to be achieved to allow sufficient
nutrient diffusion and vascularization.
PPF has also been developed for use in combination
with cell transplantation applications through the incor-
poration of cells within the material during the cross-linking
procedure. Because of the potentially non-cytocompatible
conditions that may be present during the curing reaction,
a composite material has been developed in which cells are
first encapsulated in gelatin microspheres, and these are
then included with the PPF during cross-linking ( Payne
et al. , 2002a , b). It has been shown that this encapsula-
tion procedure enhances the viability, proliferation, and
osteogenic differentiation of rat-marrow stromal cells (as
compared to nonencapsulated cells) when cultured on
cross-linking PPF in vitro ( Payne et al. ,2002a ,b).

Prevascularization

The major obstacle in the development of large 3D
transplants such as liver is nutrient diffusion limitation,
because cells will not survive farther than a few hundred
microns from the nutrient supply. Although the scaffold
can be vascularized postimplantation, the rate of vascu-
larization is usually insufficient to prevent cell death
inside the scaffold. In this case, prevascularization of
the scaffold may be necessary to allow the ingrowth
of fibrovascular tissue or uncommitted vascular tissue
such as periosteum (layer of connective tissue covering
bone) before cell seeding by injection ( Fig. 7.1.4-1 C).
The prevascularized scaffold will provide a substrate for
cell attachment, growth, and function. The extent of
prevascularization has to be optimized to allow sufficient
nutrient diffusion as well as enough space for cell seeding
and tissue growth ( Mikos et al. , 1993c ).
Some complex osseous defects created by bone tumor
removal or extensive tissue damage exceed the critical
size for normal healing and require a large transplant
to restore function. A novel strategy is to prefabricate
vascularized bone flaps by implanting a mold containing
bioresorbable polymers with osteoinductive elements
onto a periosteal site remote from the defect where
prevascularization and ectopic bone formation can occur
over a period of time as the scaffold degrades (Thomson
et al., 1999). The created autologous bone can then be
transplanted to the defect site where vascular supply can
be attached via microsurgery to existing vessels.

Delivery of bioactive molecules

Cellular activities can be further modulated by various
soluble bioactive molecules such as DNA, cytokines,
growth factors, hormones, angiogenic factors, or immu-
nosuppresant drugs ( Babensee et al. , 2000; Holland and
Mikos, 2003; Kasper and Mikos, 2003 ). For instance,
bone morphogenetic proteins (BMPs) have been identi-
fied as a family of growth factors that regulate differen-
tiation of bone cells ( Ripamonti and Duneas, 1996 ).
Controlled local delivery of these tissue inductive factors
to transplanted and regenerated cells is often desirable.
This has led to the concept of incorporation of bioactive
molecules into scaffolds for implantation. These factors
can be bound into polymer matrix during scaffold pro-
cessing ( Behravesh et al. , 1999; Shin et al. , 2003a )
( Fig. 7.1.4-2 A). Alternatively, bioresorbable microparti-
cles or nanoparticles loaded with these molecules can be
impregnated into the substrates ( Hedberg et al. , 2002;
Holland et al. , 2003; Lu et al. , 2000 )( Fig. 7.1.4-2 B). By
incorporating BMPs or other osteogenic molecules into
the injectable paste, PPF can also serve as a delivery
vehicle for bone growth factors to induce a bone-
regeneration cascade ( Hedberg et al. , 2002 ). The release
of bioactive molecules in vivo is governed by both dif-
fusion and polymer degradation ( Hedberg et al. , 2002;
Holland et al. , 2003 ). In addition, if the molecules are

In situ polymerization

Injectable, in situ polymerizable, bioresorbable materials
can be utilized to fill defects of any size and shape with
minimal surgical intervention ( Fig. 7.1.4-1 D). For instance,
PPF has been developed as an injectable bone cement that
hardens within 10 to 15 minutes under physiological con-
ditions. These materials do not require prefabrication but
must meet additional requirements since polymerization or
cross-linking reactions occur in vivo. All reagents and
products

must be biocompatible, and the reaction