The three primary components of the afferent vascular system to bone tissue are the principal nutrient artery, the metaphyseal arteries, and the periosteal arterioles. The nutrient and metaphyseal arteries together compose the medullary arterial system which is the major afferent supply nourishing about the inner two thirds of the bone cortex. The periosteal vessels enter the cortex mainly at sites of fascial and muscle attachment and appear to supply the outer third of the bone diaphysis [33–38]. Cortical circulation usually flows in a centrifugal direction. In the diaphysis the inner cortical layers are drained through venous channels, the periosteal layers directly by periosteal capillaries. In case of damage to the medullary system following trauma or operation a compensatory flow reversal occurs to some extent [33, 38–42].

As a result of bone fragmentation and displacement of fragments, periosteal, intracortical and endosteal vessels are ruptured [13, 35–37, 41, 43]. At each fracture line all intracortical vessels are disrupted due to the direct damaging of its surrounding osteons. Major displacement of the fracture fragments may disrupt larger vessels like the nutrient artery, the central artery or its intramedullary branches. This disruption of the medullary blood supply in turn leads to avascularity and devitalization of a large amount of the bone cortex. The stripping of the periosteum with its vascular network during injury is of particular importance, because disruption of the periosteum may be severe or total between fragments leaving smaller fragments completely devascularized.

The surgical approach to the fracture leads to an additional considerable vascular damage to the bone tissue by the soft tissue retraction. Additional damage is added by subperiosteal exposure that leads to more damage compared to careful epiperiosteal exposure [44]. Fragment manipulation by reduction clamps and the plates itself result in further damage to the blood supply of bone. Complete visualization of the fracture area is needed neither for fracture reduction nor for positioning of the plate and insertion of the screws [15, 45, 46].

In conventional plating technique some amount of contact between plate and bone is needed for stability reasons to allow load transmission by friction at the interface. The axial screw force generated by tightening the screws and the compressive strength of cortical bone gives the minimum area required for load transfer. Shaping the plate to the bone surface as exactly as possible is mandatory not to be faced with the problem of secondary fracture dislocation when the fragment is pulled towards the plate by tightening the screws. The biological disadvantage of the conventional contact plating concept is the appearance of a relatively large zone of blood supply disturbance directly underneath the implant (Fig. 29.1a, b). This deficiency of perfusion is caused by direct compression of the periosteal vascular network under the plate and leads to necrosis of cortex adjacent to the plate [47–55]. Dead bone can only be revitalized by removal and replacement (creeping substitution), a biological process which takes a long time to be completed. During the recovery of the blood supply, a temporary porosis of the bone is observed as a result of the tremendous intracortical remodeling. The remodeling activity starts at the boundary between vital and initially devascularized bone and is usually directed towards the implant (Fig. 29.2a, b). It is accepted that necrotic tissue disposes to and sustains infection [56]. The recovery of the original bone structure and vitality generally takes more than one year. In the past, many authors have tried to explain this temporary bone porosity as a functional adaptation of the bone structure to the unloading effect of the plate according to Wolff’s law [57–69]. The newer generations of plates (limited contact, no contact implants) decrease the amount of devascularization of cortical bone due to a reduction or the complete absence of implant-bone contact.

Fracture healing is the recovery of the biological and mechanical integrity of the osseous tissue, i.e., return of the prefracture tissue vitality and structure as well as the prefracture stiffness and strength of the injured bone segment [70]. The amount of stability achieved by implants is the mechanical input for the biological response of bone healing. Beside the injury itself, the healing process additionally is modulated by the additional surgical damage to the bone and surrounding soft tissue envelope during the process of reduction and fixation [9, 25, 71, 72]. In plate osteosynthesis the importance of the amount of mechanical stability to achieve direct bone healing was overestimated for a long time. Forcing precise reduction to improve the postoperative radiological appearance was likely to be linked to additional and sometimes extensive surgical trauma with stripping and denuding of bone fragments. Because dead bone is unable to heal some of the possible complications such as deep infection, non-union, delayed union, and refracture have to be attributed to the iatrogenic surgical tissue damage during the operative procedure. Radiographically and histologically different healing patterns can be differentiated depending on the local mechanical environment [3, 6, 73].

Absolute stability is present when the fracture is stabilized by a stiff implant which maintains the fracture reduction with no or minimal displacements occurring under functional loading. As a biological consequence primary bone healing without radiographically visible callus formation occurs. It can be assumed that fragment end necrosis induces internal remodeling of the bone which repairs the fracture by the effect of crossing osteons.

Flexible fixation allows the fracture fragments to displace in relation to each other when load is applied. The external load results only in reversible deformation of the splint. After unloading, the fracture fragments move back into their former relative position. When the load results in an irreversible deformation of the splint, the fragments remain permanently displaced. Such a situation with plastic deformation of the implant is called unstable fixation. All fracture fixation devices possess different degrees of implant stiffness and lead to fixations of gradually differing flexibility depending on how they are applied and loaded. It appears likely that some flexibility of fixation is the most important mechanism triggering and inducing callus [74].

In bone healing, the strain conditions of the involved tissues have to be taken into account when judging under which condition bridging by bone formation will occur or a non-union develops (Table 29.2). The comminution reduces the strain magnitude of the interfragmentary tissues in each gap for a given amount of overall displacement, thus allowing its safe differentiation and ossification. On the other side, a small gap with some instability still present increases the strain of the interfragmentary repair tissue with inherent risk for non-union [75].

There are two basic mechanical principles how a fractured bone can be stabilized: interfragmentary compression and splinting. Interfragmentary compression functions by the elastic preload of both, the bone and the implant. Thereby, the plate is loaded in tension and the bone in compression creating high amounts of friction between the bone fragments. Interfragmentary compression can either be static, i.e., induced as a result of a pretensioned implant, or dynamic, i.e., generated by means of the functional load allowing coaptation of the fragments along a non locked internal or external splint. Static interfragmentary compression can be accomplished only in case of at least a partial bony contact between the main fragments. Additionally, interfragmentary compression is very sensitive to minimum amounts of motion-induced bone resorption, diminishing the preload of both, the implant and the bone, with consecutive loss of stability [76, 77].

Splinting consists of the connection of an implant to a broken bone. The stability of this composite system depends on the stiffness of the splint itself, the quality of coupling between the splint and the bone, and the presence or absence of fracture comminution and bone defects. Depending on the localization of the implant, splinting can either be external or internal. Internal splints can be positioned inside or outside the medullary cavity. Depending on the mechanical use of the implant, a splint can be either gliding (nonlocked or dynamically locked) or nongliding (statically locked). The plate design itself does not define its later mechanical function; it can be used for splinting and/or for compression osteosynthesis.

The Dynamic Compression Plate (DCP) was introduced in 1969 [8]. The idea of this plate was to enhance stability of fixation by interfragmentary compression and load transfer by friction (Fig. 29.3a-c). To improve periosteal vascularity underneath a plate, in a first step the plates were undercut as far as safe application of the plate screws would allow without exceeding the compressive strength of the underlying cortical bone. The Limited Contact Dynamic Compression Plate (LC-DCP) which is still in clinical use was the first implant modified to preserve the bone circulation [12, 78, 79]. Nevertheless, some contact between bone and implant is still needed to allow load transfer by a friction force at the interface created by tightening the screws.

The next step consisted in minimizing the plate-bone contact to isolated points only. With the Point Contact Fixator (PC-Fix) the plate is not compressed towards the bone. The isolated contact points between the plate holes serve only to hold the plate at a distinct distance from the bone surface as soon as the screw heads engage in a conical non-threaded plate hole [80]. With the PC-Fix the load transmission is based on the partial interlocking of the screw heads in the plate holes. Thus, the amount of contact of the implant remains very small and without any adverse biological or mechanical effects to the underlying bone (Fig. 29.4).

Nowadays, conventional plating is increasingly replaced by using internal fixators. Internal fixators are “plates” (splints) with completely locked screw heads. These implants are not pressed onto the bone and do not need any contact between implant and bone. Further advantages are the possible reduction of the screw length to monocortical dimension with the advantage not to need screw length measurements and the possibility of using self-drilling and self-tapping screws.

The Less Invasive Stabilization System (LISS) was the first internal fixator of the AO-ASIF conceived for the meta- and epiphyseal regions of the distal femur and proximal tibia (Fig. 29.5). Its shape conforms to the anatomical contours of the specific area of the bone and is designed for application via a minimally invasive submuscular approach. First step is the anatomic reconstruction of the articular component followed by the restoration of the correct bone axis in all planes using the femoral distractor [81, 82].

A further development is the Locking Compression Plate System (LCP) which offers the advantage that the surgeon can choose during the operation whether to use it with conventional (non-locked) screws, with locked screws, or with a combination of both [83]. The specific design of the plate hole shows two functional elements: The first half of the hole comprises a dynamic compression unit that is intended for a standard cortical or cancellous bone screw. As in standard DC-plating technique the eccentric screw insertion allows an axial compression at the fracture site to be achieved. Additionally, the screw can be angulated with respect to the longitudinal and transverse plate axis. The second half of the hole is threaded and conically shaped permitting the locking of the special locking head screws (Fig. 29.6a-c). The conical shape of the threaded screw head stops the tightening well before the thread within bone sustains critical loads. This in addition eliminates plastic deformation of the screw while tightening and equally excludes the reported high incidence of screw failures during insertion of conventional screws. No contact between bone and implant is needed for load transmission. The angular stability of the screws results in a lower incidence of screw loosening and secondary displacement of the fracture fragments. In addition, the protection of the periosteal blood supply is superior and in subcutaneous and submuscular plating techniques there is no need for accurate contouring of the plate [84–86].

Due to the surgical needs for plate-screw systems allowing locking of the screws in different directions other systems were developed. The plate hole of the Variable Angle Locking Compression Plate has four columns of threads providing four points of locking between the Variable Angle LCP and the variable angle screw, forming a fixed-angle construct at the desired screw angle (Fig. 29.7a, b). Screws can be angulated anywhere within a 30° cone around the central axis of the plate hole (Fig. 29.8a, b). The variable angle LCP combi hole combines a dynamic compression hole with a variable angle locking screw hole and provides flexibility of selecting either axial compression or variable angle locking in the same plate hole.

Another technical solution is found with the noncontact bridging plate (NCB) technology allowing polyaxial screw placement with screw locking achieved through the use of locking caps that are threaded into the plate holes. Nevertheless, due to the fact that the locking of screw with the locking cap is performed after screw tightening the biological advantage of a no contact system can be abolished (Fig. 29.9). On the biomechanical side all this variable angle screw locking system are as safe as the plate systems with fixed angle locking screws [87–89].

In a conventional plating technique, insertion and tightening of the plate screw generates an axial screw force (F_a), which compresses the plate onto the bone surface. This compression leads to a friction force (F_f) at the interface bone-implant. The friction force is proportional to the amount of compression with the specific bone-implant coefficient of friction (ρ) as a constant of proportionality. Under functional loading, a plate osteosynthesis is loaded mainly in bending and in axial load. These loading patterns tend to displace the plate with regard to the bone. As soon as the external load (F_e) exceeds the amount of friction force installed at the plate-bone interface, the plate will slip on the bone (Fig. 29.10). Under stable conditions, the screw is mainly loaded in tension. Once the plate starts slipping on the bone surface an additional bending moment at the screw head-screw shaft junction is present [77]. The advantage of the conventional plating is the possibility to angulate the screws with respect to the plate surface (Fig. 29.11). The disadvantage is that with time the axial screw force is diminished due to bone remodelling around the screw threads leading to a corresponding decrease of the friction force at the bone-implant interface. In addition, in osteoporotic bone the maximum screw force that can be obtained by screw tightening can be very low from the beginning resulting in a low friction force and later instability.

In contrast to the conventional plating technique, in “locked screw head” plating no contact between implant and bone is needed for load transfer from one main fragment to the other. Therefore, no friction force is generated by axial screw force at the interface bone-implant to withstand the displacement forces. Mechanically this so-called internal fixator is defined as a construct in which the screws are the principal load-transferring elements from the main bone fragments to the implant. In such a construct the screws are firmly locked to the internal fixator to allow for moment and force transfer (Fig. 29.12). Thus, the longitudinal displacement forces acting on the construct are directly transferred from the bone to the implant by bending and shear across the screw neck [45, 80]. The advantage of the locked screw head is the change of the mechanical loading condition of the screw, which now is mainly in bending and less in axial pull-out. The disadvantage is the complete loss of the surgical feeling of screw tightening. Even in weak bone, the threaded screw head engages firmly inside the plate giving false information about the holding of the screw inside the bone. The former disadvantage of the fixed angle screw direction inside the LCP is resolved with the newer generation of Variable Angle Locking Compression Plates.

In osteoporotic bone the commonly recommended fixation concepts fail, because bone quality is poor. Holding power of each screw in conventional plate osteosynthesis is decreased leading to the problem of early pullout of screws and secondary fracture displacement. A possibility to enhance the quality of fixation is the use of bone cement. Bone cement can be used around blade plates to diminish the danger for cut-out and around the intramedullary part of the screws to enhance the holding power. In such a case, two screws should be angulated to each other and the intramedullary space in between should be filled with bone cement (Fig. 29.13). After polymerization the screws are tightened. The pairwise obliquely inserted screws give a very stable fixation of the plate.

In osteoporotic bone fracture fixation using the internal fixator concept seems to be advantageous, because stability of fixation does not rely on an axial screw force creating a friction force at the implant-bone interface, but on locked screws loaded mainly in bending during functional loading.

Another possibility to increase the holding of screws in osteoporotic bone is the use of “schuhli nuts”. Schuhli nuts are placed underneath the plate and have an identical thread like the conventional cortical screws. Tightening the screw locks it within a plate hole creating an angular stability of the screw-plate-schuhli construct with enhanced holding power (Fig. 29.14). The advantage of the schuhli nuts is avoiding the potential adverse effects of bone cement on fracture healing with extravasation and thermal necrosis [90–92]. But, the fiddling factor of the schuhli nuts is more important than with the use of a locking screw head system (LCP, LISS).

Most plates used for osteosynthesis are still metallic implants (stainless steel or titanium). The availability of the material and more important the excellent mechanical and biological properties of metals are the main reasons for its widespread use in internal fixation. Metals offer on the mechanical side high stiffness and strength to withstand deformation or fatigue failure, sufficient ductility to allow shaping of the implant, corrosion resistance and on the biological side good tissue compatibility without localized toxic reactions [93–96].