Klaus Louis Gerlach and Uwe Eckelt

Different chapters of this book describe applications of various mini/microplate and -screw systems that guarantee in nearly all cases a secure and successful bone healing of osteotomized or fractured bone segments, in mandibular and midfacial fractures as well as in orthognathic procedures. Plates and screws are generally made of titanium, and until now these have been regarded as the “gold standard.”

Some disadvantages have become apparent, however, such as interference with later diagnostic or therapeutic radiological investigations (e. g., computed tomography, magnetic resonance imaging), under-the-skin palpable plates, growth inhibitions of craniofacial bones in infants followed by passive migration, loosening of screws, hot and cold irritabilities, and, not least, secondary intervention to remove osteosynthesis materials. These considerations have led to the development of biodegradable osteosynthesis materials as possibilities for more than three decades (Kulkarni et al., 1971; Champy et al., 1993; Gerlach, 2000).

The major advantages of biodegradable osteosynthesis devices are that functional stress is gradually transferred to the bone as it remodels and matures, and the material hardly ever requires surgical removal.

Among the various resorbable polymers used for this purpose, semicrystalline poly-L-lactide (PLLA), amorphous poly-D,L-lactide (PDLLA), polyglycolide (PGA) and their copolymers (P[L/LD]LA and PLGA) are the most important and commonly used.

The mechanical characteristics, the biocompatibility, and the duration of biodegradation vary among the different systems available. These depend upon the polymers or copolymers used, their amount of crystallinity, and their molecular weight. The methods used to produce and form the materials, for example self-reinforced techniques (SR), the sterilization process, the amount of materials used, and the perfusion of tissues overlying the material once implanted also differ according to the system.

Polylactic acid (PLA), which is the most often used, has two enantiomeric forms, L-lactic acid and D-lactic acid. Whereas in L-lactide identical polymer chains can be tightly packed, which makes these polymers partially crystalline, adding D-isomers into an L-isomer-based polymerization system balances the tightening of the packing and, depending of the amount of D-lactide, reduces the crystallinity or results in an amorphous copolymer. Although both PLLA and PGA are highly cristalline, their co-polymer, for example with a PLLA:PGA percentage ratio of 82:18, is amorphous.

The mechanical characteristics are especially improved by an increase of the molecular weight, but on the other hand these properties prolong the duration of biodegradation.

At present numerous different systems are clinically available. These are, however, characterized by the polymer used or their respective polymer compositions; also by differing physical properties, biological compatibilities, and absorption times resulting in persistence lasting between 1 year and more than 5 years (Buijs et al., 2006). Moreover, the primary and continuous tension or bending strength, and especially the stiffness, are much lower with all available systems compared with conventional materials made from stainless steel or titanium (Daniels et al., 1992). The application is therefore especially recommended for the stabilization of sections of the face that are not strongly load-bearing (midface and cranium), and exceptionally in growing patients. The following discussion focuses particularly on our own experiences with a system made from poly-D,L-lactide.

The source material is 50: 50 poly-D,L-lactic acid (PDLLA), from which pins, micro and mini osteosynthesis screws are produced using an injection molding process. The osteosynthesis plates and meshes are pressed and milled from granulated polymer. Prior to sterilization, the plates have a molecular weight (M_w) of 200000g/mol, a melting point of 120°C (248°F), a tensile strength of approximately 60 MPa, and a bending strength of 120 MPa. The material is gamma-sterilized with an overall dose of 25 kGy.

After in-vivo examination, a slight decrease of the bending strength of 15 % was observed within 6 weeks postimplantation; thereafter a rapid falling off of the mechanical properties occurred (Heidemann et al., 2003a).