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EXTERNAL SKELETAL FIXATION

PRINCIPLES OF BONE HEALING AND BIOMECHANICS OF EXTERNAL SKELETAL FIXATION Ross H . Palmer, DVM, MS, Donald A. Hulse, DVM, William A. Hyman, ScD, PE, and Darcie R. Palmer, DVM

Although gaining in popularity among veterinary orthopedists in recent years, external skeletal fixation has a reputation for causing excessive complications, including premature loosening of fixation pins, implant breakage, and pin tract sepsis. These contribute to patient morbidity through poor limb use, pain, loss of fracture reduction, delayed healing, and nonunioris. Veterinary surgeons therefore need a thorough understanding of basic biomechanical principles of external fixation to reduce patient morbidity. ORIGIN OF LOADS APPLIED TO FIXATORS

Long bones of the skeleton are subjected to forces generated by weight bearing, muscle contraction, and physieal activity in the reha~ biIitation period. The force of weight bearing occurs as the foot makes contact with the ground. Simultaneously the ground responds with an equal but opposite reaction referred to as ground reaction force .7 The magnitude of the ground reaction force is proportional to the relative From the Santa Cruz Veterinary Hospital, Santa Cruz, California (RHP); the Department of Small Animal Surgery, College of Veterinary Medicine (DAH, DRP), and the BioEngineering Program, College of Engineering (WAH), Texas A&M UniverSity, College Station, Texas

VETERINARY CLINICS OF NORTH AMERICA: SMALL ANIMAL PRACTICE VOLUME 22· NUMBER 1 • JANUARY 1992

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amount of body weight carried by the foot at impact and the body acceleration (force = mass x acceleration).l1 Because of deceleration, the ground reaction force may increase to more than five times body weight during jumping. 40 The ground reaction force causes axial compression and bending and torsional moments in bone. 27 In a fractured bone stabilized with an external fixator, these forces and moments are shared in varying amounts with the components of the fixator. BASIC BIOMECHANICS

Physiologic forces applied to bone arise from joint interfaces and muscular attachments and are uniaxial (tension or compression). These forces, however, can give rise to torsional and bending moments within the long bones and within the external fixator. A bending moment occurs when a force is applied to an object such that it causes the object to bend about an axis (Fig. 1) . 27 A torsional moment occurs when an applied force causes the object to twist about an axis (Fig. 2).27 The magnitude of a bending or torsional moment is the product of the applied force times the moment arm (lever arm) over which it acts. 27 The moment arm is equal to the perpendicular distance from the line of force application to point at which the moment is calculated (Fig. 3). Joint and muscle forces also result in shear stresses within the bone. Fracture healing and return of normal locomotor function depend on counteracting the forces, moments, and shear stresses acting at the fracture site. Functionally the most important mechanical properties of a fixator-bone composite are strength and stiffness. These properties are best understood by examining the behavior of a given object under

Figure 1. Uniaxial forces applied eccentrically to a bone column may create a bending moment about its central axis. (Example: compression applied to femoral head, which is positioned 10 eccentric to central axis of the femur, creates bending moment about the femoral central axis) .

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Figure 2. Uniaxial forces applied eccentrically to a bone column may create a torsional moment about the column's central axis. (Example: tensile forces exerted by the iliopsoas muscle at its insertion on the lesser trochanter of femur create a torsional moment about the femoral central axis).

loading. Application of known load causes a measurable deformation of the object and can be plotted on a load-deformation curve (Fig. 4). For many materials, load and deformation are linearly related in the initial portion of the curve. In most cases, the initial portion of the curve, where load and deformation are linearly related, represents elastic deformation, i.e., the object returns to its original shape when the applied load is removed. For "ductile" materials (316 L stainless steel), the elastic limit of the object is indicated by the yield point. As

t

MOMENT ARM

Figure 3. Application of uniaxial compressive force to wrench handle (eccentric to central axis of bolt) exerts a torsional moment about the center of the bolt. Magnitude of torsional moment equals applied force x moment arm. The moment arm is the distance from the applied force (arrow at end of wrench handle) to the central axis of the bolt. The magnitude of the torsional moment can be increased by application of greater force or by increasing the moment arm.

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ULTIMATE STRENGTH PLASTIC DEFORMATION

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Figure 4. Load-deformation curve. Application of known load to an object causes measurable deformation. A represents the starting point. 8 is the yield point beyond which "ductile" materials undergo plastic deformation (deformation remains after removal of load). C is the ultimate strength where breakage occurs.

load exceeds the yield point, the object undergoes plastic deformation, Le., permanent deformation remains after the object is unloaded. If the load is progressively increased, the object will break; this point defines its ultimate strength. "Brittle" materials (glass) do not exhibit plastic deformation, and the ultimate strength may be within the linear region of the curve. The stiffness of the object is generally determined by the slope of the curve in the linear elastic phase. The steeper the slope, the stiffer the object. One must distinguish between material properties and structural properties of objects (bone, implants). Material properties of an object are independent of its dimensions and depend only on material from which the object is constructed. 27 Structural properties are dependent on dimensions.27 An example is the comparison of two fixation pins made of the same stainless steel material but one pin being twice the diameter of the other. A force-deformation curve can be generated for the two pins (Fig. SA). In this comparison of structural properties, the larger pin is stiffer and has a greater yield point and ultimate failure point. To compare material properties of the two pins, local forces, which are produced within the pins during force-induced deformation, must be determined. Local force magnitude for a specified area of each pin is referred to as internal stress; local deformation for a specified area is referred to as internal strain . Internal stresses have the dimensions of force/cross-sectional area, and internal strains are equal to the change in length/original length (axial strain) or a measure of change in angles (shear strain). A stress/strain curve depicted for the two pins of identical

PRINCIPLES OF BONE HEALING AND BIOMECHANICS

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STRAIN B Figure 5. A, Load·deformation curve for two different-sized stainless steel pins. Structural properties are dependent upon dimension, and the large pin has greater stiffness and ultimate strength than the small pin. B, Stress-strain curve for the same two pins. Material properties are independent of dimension. This curve shows that the pins have identical material properties.

material but differing dimensions illustrates they have identical material properties (Fig. 5B).

EFFECT OF FIXATION VARIABLES ON FIXATOR STIFFNESS AND STRENGTH

Load-sharing between the fractured bone and the implant, number and size of fixation pins, pin structural design, pin position, angling of pins, and frame configuration directly affect stiffness and strength of the fixator-bone composite.

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Load-Sharing and Non-Load-Sharing Fixation

When a fixator is applied for fracture stabilization, a varying degree of load sharing is experienced by the fixator and by the bone column.! When a transverse fracture is reduced and the bone segments compressed, much of the ground reaction force is transmitted axially through the bony column (Fig. 6A) . Therefore loading of the fixator and the pin-bone interfaces is minimized in instances when the fractured ends of the bone column can be compressed. Highly comminuted fractures that cannot be anatomically reconstructed, segmental bone resections, and limb-lengthening procedures represent buttress situations in which none of the ground reaction force is transmitted through the bone column; instead all of the load is transmitted from bone segment to bone segment through the pin-bone interfaces and fixator components (Fig. 6B). When bones with comminuted or oblique fractures can be anatomically reconstructed and interfragmentary compression achieved with full cerclage wire or lag screws, a fixator shares the

Figure 6. A, Ideal load-sharing fixation. Transverse fracture has been anatomically reduced; much of the ground reaction force is transmitted axially through bone column. Relatively little loading of fixator components and pin-bone interfaces occurs. B, Nonload-sharing fixation. None of the ground reaction force is transmitted through bone column; instead, all of the load is transmitted from bone segment to bone segment through the pin-bone interfaces and fixator components. Illustration continued on opposite page

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Figure 6 (Continued). C, Partial load-sharing fixation . Oblique fracture has been reduced, and the bone column reconstructed using four fuJI-cerclage wires for interfragmentary compression. The fixator is being used in a neutralization mode; it is neither acting to compress the bone segments nor hold them apart. In this instance, there is some load sharing between the fixator and bone column.

axially transmitted ground reaction force with the reconstructed bone column (Fig. 6C). In this situation, the fixator is being used in neutralization fashion. Pin Number

More fixation pins increase stiffness of the fixator-bone composite. In a study using bilateral frame configurations, stiffness increased with increased pin number up to four pins per bone segment; use of more than four pins per bone segment did not significantly increase fixatorbone composite stiffness.13 Interestingly in buttress (non-load-sharing) situations, the connecting rod of unilateral fixator frames is the weakest component against axial compressive loads, and failure occurs by plastic deformation of the connecting rod at the level of the osteotomy.16 For this reason, increasing from two to four pins per bone fragment did not significantly increase the axial stiffness of unilateral frames with single connecting rods. IO When two connecting rods were used with a four-pin unilateral frame (buttress situation), failure occurred by plastic deformation of the fixation pins and connecting rods. 16 Accordingly increasing from two to four fixation pins per bone fragment approximately doubled the axial compressive stiffness of unilateral frames with two connecting rods. 10

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Half-Pins Versus Full-Pins

Frames consisting entirely of full-pins (bilateral-uniplanar frames) are stronger against axial compression, transverse loading of the osteotomy, and torsion than frames made entirely of half-pins (unilateraluniplanar frames) . 10, 16 The most proximal and distal full-pins of a bilateral frame are easily placed, but placement of full-pins adjacent to the fracture is more technically difficult. In such a situation, half-pins placed through each pin clamp adjacent to the fracture (instead of fullpins) caused an approximately 25% loss of axial stiffness. 1o The effect of half-pins versus full-pins on bending and torsional stiffness in such situations has not been thoroughly investigated. Pin Diameter

Pin diameter has a tremendous effect on stiffness of the pin itself and therefore on the fixator-bone composite. Because stiffness is directly related to the fourth power of radius, increasing the diameter of a pin from 2 mm to 4 mm increases its stiffness by a factor of 16.13 Experimentally increasing the pin diameter from 2.5 mm to 4.0 mm more than doubled the overall stiffness of a variety of bilateral fixator frames regardless of their specific design. 10 Pin DeSign

The pin-bone interface is the most highly stressed portion of a fixator-bone composite. 13 High pin-bone interface stress leads to bone resorption and premature loosening of fixation pins, the most common complication of external fixation. Increasing numbers of threaded and non threaded fixation pin designs are available to the veterinary orthopedist (Fig. 7). Threaded pins have greater holding power than nonthreaded pins and are used to decrease the incidence of premature loosening and the associated morbidity.3,29 Many threaded pin designs are available. There are completely threaded pins, end-threaded pins, and central-threaded pins. End-threaded pins are used in arrangements requiring half-pins (type I and type III frames), whereas central-threaded pins are used in arrangements requiring full-pins (type II and type III frames). Completely threaded pins can be used either as half-pins or as full-pins. Because pin stiffness is related to the fourth power of the pin radius, threads that are cut into a standard non threaded pin decrease its stiffness; therefore completely threaded pins are less stiff than nonthreaded pins of similar outer diameter.6 Some surgeons attempt to overcome the loss of stiffness of completely threaded pins by using end-threaded fixation pins, but the stiffness of end-threaded pin-bone composites is dependent on the length of thread and the amount of

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Figure 7. A, Fixation pins implanted as half-pins. B, Pins implanted as full-pins. Pin no. 1 is a nonthreaded Steinmann pin used both as a half- and full-pin. Pin no. 2 is a completely threaded Steinmann pin used both as a half- and a full-pin. Pin no. 3 is a two-cortex endthreaded pin (threads engage two cortices) used as a half-pin. Pin no. 4 is a one-cortex end-threaded pin (threads engage only the far cortex) used as a half-pin. Pin no. 5 is a two-cortex end-threaded pin with a positive thread profile (outer thread diameter greater than shaft diameter) used as a half-pin. Pin no. 6 is a centrally threaded pin with a positive thread profile used as a full-pin.

thread engaged in the bone. 31 The threads of some end-threaded pin designs engage both the near and the far cortex of bone (double-cortex, end-threaded pins) leaving some of the threads and the threadednonthreaded shaft junction outside the bone (see Fig. 7, pin A-3). The threads of other end-threaded pin designs engage only the far cortex of bone (single-cortex, end-threaded pin) such that the threadednonthreaded shaft junction is positioned within the intramedullary canal (see Fig. 7, pin A-4). Single-cortex, end-threaded pin-bone composites are stiffer than equal diameter double-cortex, end-threaded pin-bone composites in which some thread is exposed. 6 Similarly single-cortex, end-threaded pin-bone composites are stiffer than equal diameter completely threaded pin-bone composites. 6 There is little difference in pin-bone composite stiffness between completely threaded and double-cortex, end-threaded pins of similar diameter.6 Moreover because of stress-concentration at the threaded-non threaded shaft junction in end-threaded pins, double-cortex, end-threaded pins become less stiff than completely threaded pins after they are subjected to 10 degrees of deflection. 6

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Pins with an outer thread diameter greater than the shaft diameter are said to have positive thread profiles. Such pins have the advantage of good holding power without the disadvantage of loss of stiffness (that pins with "cut" threads have) (see Fig. 7, pins A-5 and B-6).

Pin Separation

The bending stiffness of the connecting rods is inversely related to length to the third power. The working length of the connecting rod can be minimized and stiffness maximized by placing pins in close proximity to the fracture. In midshaft fractures, the remaining pins should be evenly distributed throughout each bone fragment. Experimentally doubling the distance from the most proximal to the most distal pin in each bone fragment (44 to 90 mm) caused little change in stiffness against bending moments applied in the same plane as the fixator but almost doubled the stiffness against bending moments applied in a plane perpendicular to uniplanar (unilateral and bilateral) fixator frames.5 Mathematical modeling suggests that in proximal or distal fractures, pins should be maximally separated in the small fragment; in the large fragment, the distance between pins should be approximately twice the inter-pin distance of the small fragment. 18

Angled Versus Parallel Fixation Pins

Angling the most proximal and distal pins approximately 20 degrees inward (toward the fracture) maximizes fixator stiffness.2o When recent studies that used similar osteotomy models are compared, it is noted that the incidence of severe pin loosening was much greater in the model where parallel pins (versus angled pins) were used. 21.29

Number of Connecting Rods

In static axial compression tests, failure of four-pin (single rod) unilateral fixator configurations occurred by bending (plastic deformation) of the connecting rod at the level of the osteotomy.16 Use of a second connecting rod with unilateral frames approximately doubled the strength against axial compression, and failure occurred by plastic bending of the connecting bars and fixation pins. 16 Resistance to transverse loading at the osteotomy site was also approximately doubled, but there was little effect on torsional strength. Addition of a second connecting bar to a unilateral frame increased bending stiffness by approximately 20% both in the plane of the fixator frame and in the perpendicular plane. 5

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Frame Configuration

Static strength evaluation of different configurations has shown unilateral (type Ia), unilateral-biplanar (type Ib), bilateral-uniplanar (type II), and bilateral-biplanar (type III) frames to be successively stronger in resisting axial compression and torsion. 16 The same relationships are true for resistance to transverse load applied at the osteotomy, with the exception that unilateral-biplanar . (type Ib) frames were stronger than bilateral-uniplanar (type II) frames. 16 Connecting Rod to Bone Distance

Stiffness of fixation pins is inversely proportional to their length to the third power. Therefore placement of the connecting bar progressively closer to the bone shortens the working length of fixation pins and increases their stiffness. Experimentally as the connecting rod to bone distance decreased (from 80 mm to 25 mm) with a four-pin unilateral fixator, resistance to bending in the plane of the fixator was unchanged, while resistance to bending in the plane perpendicular to the fixator frame increased over 200%.5 Mathematical models have also shown that pin length has no significant effect on stiffness against bending moments applied in the plane of the fixator but is significant in resisting bending moments applied perpendicular to the plane of the fixator.14 The effect of specific fixation variables on fixator stiffness and strength are reasonably well defined, and efforts by clinical orthopedists to maximize stiffness and strength have been effective in reducing complication rates to a reasonable level during the past decade. It is important to realize, however, that the ideal stiffness at the fracture line for maximized fracture healing has not been determined and will likely be an area for clinical studies in the coming years. MECHANISMS OF IMPLANT FAILURE Premature Loosening of Fixation Pins

Premature loosening of fixation pins is the major complication with external fixators.1, 2, 3 Pin loosening can contribute to patient morbidity by causing severe pain, poor limb function, fracture instability, delayed unions, and nonunions. It is considered the most significant feature in the development of pin tract osteomyelitis. 2, 3 Implant loosening is the result of a complex series of interrelated mechanical and biological events.9, 25 Variable degrees of microstructural bone trauma caused by implant insertion is the initial event in implant 100sening. 38, 39 In vitro study has shown that cyclic implant loading may propagate microcracks and cause further implant 100sening. 38 Mechanical, vascular, and ther-

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mal damage cause a variable-sized rim of dead bone surrounding fixation pins soon after insertion.25, 34-36 The fate of the interfacial tissues between the implant and the live bone is heavily influenced by the local strain environment. 25, 30, 34, 35, 37 When interfacial tissue strain is small (less than 2%), undifferentiated cells develop osteogenic activity and maintain a stable pin-bone interface. When interfacial tissue strain is high (more than 2%), cellular differentiation favors osteoclasis and fibrogenesis or chondrogenesis causing pin loosening.25 Osteoclastic bone resorption effectively reduces interfacial tissue strain by enlarging the space between the pin and the bone (strain = change in length/ original length). This enlarged hole loosens the grip of the pin on the bone. Once loose, physiologic loads produce more micromotion at the pin-bone interface. The result is a cycle of high interfacial tissue strain, osteoclastic resorption of bone, and more pin loosening. Pin tract sepsis or inflammation cause further pin loosening. The ideal technique for insertion of external fixation pins in small animals remains unclear and controversial. Direct insertion of fixator pins with a high-speed drill causes thermal necrosis of bone surrounding the pin and leads to premature pin loosening.1 7 A study using human cadaveric bone found the point design of pins to be more significant than drill speed in determining temperature of bone surrounding the inserted pin; trochar and spade points (commonly used in veterinary medicine) produced the highest temperatures over the greatest portion of bone. 26 In the same study, pre drilling the holes with a twist drill bit (slightly smaller than the fixator pin) before placing the fixator pins with a hand chuck effectively reduced cortical bone temperatures. 26 Studies performed in live dogs showed no difference in pin temperature when pins were inserted by slow-speed power (150 rpm) as compared with pins inserted by hand chuck after pre drilling the holes.17 Differences between the human and canine studies may relate to the relatively thin cortical bone in dogs, smaller pin diameters used in the canine study, and better heat dissipation in living tissue. Results of one canine study showed uniform holding power of fixator pins inserted by hand chuck, pre drilling, and low-speed power methods; high-speed drilling and hand drilling decreased holding power. 17 Histologic examination showed mechanical bone damage attributed to "wobble" surrounding hand chuck inserted pins, increased bone necrosis around high-speed power inserted pins, and increased inflammation around hand drill inserted pins. The technology of pin design is of ever-increasing importance in the prevention of premature loosening of fixator pins. 39 The cutting tip of an ideal pin should efficiently cut its own thread and store bone cuttings away from the cutting surface during insertion, thereby decreasing microstructural trauma and thermal necrosis. As previously mentioned, however, the trochar and spade tips commonly used in veterinary medicine do not allow for efficient cutting of bone and may contribute to thermal injury and microstructural trauma of bone. The thread design of an ideal pin should maximize holding power once inserted. Threaded pins dramatically improve holding power of

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pins over nonthreaded pins. 3, 17, 29 This is particularly true after the pins are subjected to 8 weeks of fracture/osteotomy stabilization,29 Threaded pins that engage two cortices have greater holding power than threaded pins that engage one cortex; the increase, however, is not twofold as one might expect. 17 Implants with a large outer thread diameter to inner core diameter ratio maximize holding power in cancellous boneY The role of thread pitch, thread form, and thread diameter on the holding power of fixator pins inserted into cancellous and cortical bone is incompletely defined at present.

Fatigue Failure of Fixation Pins

Breakage of fixation pins is much less common than premature loosening but does occur. 3, 19,28,29 Breakage may occur as a result of a single load in excess of a pin's ultimate strength but more commonly results from fatigue. Fatigue failure of an implant refers to breakage that occurs owing to repeated load at levels below that required to cause breakage from a single load (ultimate strength), Generally the greater the peak stress produced within a given load cycle, the fewer cycles needed to fatigue the implant. The relationship between peak stress and number of cycles to failure is represented by the fatigue life curve (or S-N curve) (Fig. 8).23 The vertical axis represents peak stress per cycle, and the horizontal axis represents the number of cycles to failure. This curve illustrates the need for careful application of implants to facilitate load-sharing between the implant and bone during the healing period and thereby reduce the peak loads on the pins, In the left-hand section of the curve, the allowable stress (point of implant failure) is reached with a low number of cycles and high loading. This can be seen in patients with non-load-sharing (buttress) fixation (Fig. 9), Similarly low cycle/high load failure may occur when a patient repeatedly jumps on the leg postoperatively, In the right-hand section of the curve, the allowable stress is such that failure occurs with low loads but a high number of loading cycles. This is observed clinically, when the patient's activity is not limited after surgery or when there is a delayed union and the implant must remain functional beyond a reasonable amount of time, Fatigue failure of completely threaded pins most commonly occurs at the pin-bone interface,3 Fatigue failure of end-threaded pins most commonly occurs at the threaded-non threaded shaft junction, 19, 23, 28 When a single-cortex, end-threaded pin (Ellis pin, Kirschner Medical Corp., Timonium, MD) was recently introduced to veterinary surgeons, it was theorized that proper placement of the failure-prone threadednonthreaded shaft junction within the intramedullary canal would shield it from stress and, thereby, decrease the incidence of fatigue failure as compared with completely threaded and double-cortex, endthreaded pins. 6 Although the incidence of failure of these single-cortex, end-threaded pins is unknown, there are recent clinical reports of their

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Peak Stress/Cycle (S)

No. of cycles to failure (N) Figure 8. Fatigue life curve (S-N curve) represents relationship between peak stress per cycle and number of cycles to failure. As peak stress increases, the number of cycles to failure decreases (left hand portion of curve). Conversely, low stress/high cycle failure may also occur (right hand portion of curve) .

failure. 19, 28 Lysis of the near-cortex of bone around these end-threaded half-pins is thought to eliminate the stress-shielding effect accomplished by placing the threaded-nonthreaded shaft junction within the intramedullary canal. 28 To provide the holding power of threaded pins without the concurrent loss of stiffness and tendency for fatigue failure, fixation pins with positive thread profiles (Turner pins, Zimmer Medical, Warsaw, IN; Centerface-and-Interface pins, IMEX Veterinary, Longview TX; EHTP and EHFP pins, Gauthier Medical, Rochester, MN) (rather than negative thread profiles of "cut" threads) have recently been introduced (see Fig. 7). These pin designs have an outer thread diameter that is greater than the core diameter and a core diameter that is approximately equal to the inner thread diameter. Until recently, the disadvantage at these pins has been cost, but several new veterinary orthopedic implant manufacturers* are selling such fixation pin designs at affordable prices. BONE HEALING WITH EXTERNAL SKELETAL FIXATION

The effectiveness of an implant in providing stability and the biological environment at the fracture surface determine the mechanism *IMEX Veterinary Inc., Longview, IX 75004, 1(800)828·IMEX; Gauthier Medical Inc., Rochester, MN 55901, 1(507)289-0761.

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Figure 9. Radiograph from canine with external fixator used as a distractor for limb lengthening. Arrows point to gap at osteotomy, which is incom· pletely bridged with callus. This represents non· load-sharing fixation such that all loads must be transmitted from 1 bone segment to the other through components of the fixator. Most proximal end·threaded pin has broken as result of high load/low cycle fatigue. Pin failed at threaded· nonthreaded junction implanted at the pin-bone interface. (From Palmer RH, Aron ON. Ellis pin complications in seven dogs. Vet Surg 19:440445:1990; with permission.)

of fracture repair. Bone union may occur by direct healing (osteonal reconstruction) or indirect healing (intermediate callus formation). Direct healing occurs through osteonal remodeling of the cortex and may be classified as primary osteonal or secondary osteonal reconstruction (Table 1).12 Primary osteonal reconstruction occurs when precise anatomic alignment of the fracture ends and absolute stability are present. The fracture line is characterized by areas of bone to bone contact and areas where small gaps of different widths are present. Primary osteonal reconstruction includes both contact healing and gap healing. Contact healing occurs in the zones of cortical bone contact and is, characterized by osteonal remodeling across the fracture plane (Fig. 10). In preparation for osteonal remodeling, cutting cones are Table 1. CLASSIFICATIONS OF BONE UNION MECHANISMS WITH EXTERNAL FIXATION Direct healing (osteonal reconstruction) Primary osteonal reconstruction Contact healing Gap healing Secondary osteonal reconstruction Indirect healing (intermediate callus formation)

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CONTACT

HEAlING

GAP

HEAliNG

Figure 10. Types of primary osteonal reconstruction under conditions of rigid stability. Contact healing occurs in zones of cortical bone contact. Cutting cones are lined with osteoclasts at the spearhead and osteoblasts along the sides so that bone resorption and formation occur simultaneously. Gap healing occurs in the small fragment gaps between contact zones (in gaps less than approximately 1 mm and with less than 2% deformation). Lamellar bone perpendicular to the fragment ends fills the gap and is later replaced by longitudinally oriented osteonal remodelling. (Note: For purposes of describing bone healing, this figure is correct; however, the plate application on the medial side of the femur is technically incorrect). (From Kaderly RE. Primary bone healing. Semin Vet Med Surg (Small Anim) 6:21-25, 1991; with permission.)

formed at the ends of the osteons nearest the fracture. Osteoclasts line the spearhead of the cutting cone for bone resorption whereas osteoblasts line the rear of the cutting cones in preparation for bone formation. Bone resorption and bone formation occur simultaneously as the cutting cones advance and cross the fracture plane from one fragment to the other at a rate of 50 to 80 f.L per day. Gap healing occurs in the small fragment gaps between contact zones. Although the bone is not in direct apposition, adequate stability is provided by the contact zones on either side of the gap. Interfragmentary deformation must be less than 2%, and the gap width must not exceed approximately 1 mm for gap healing to occur. The gap is initially filled by blood vessels and loose connective tissue. After approximately 2 weeks, the vascular supply is established and osteoblasts deposit lamellar bone in the gap between the fragment ends. The new lamellar bone fills the gap but is oriented perpendicular to the fragment ends (Fig. 10). Initially there is poor connection between the new bone and existing bone of the fragment ends, which makes this area mechanically inferior. In 3 to 4 weeks, cutting cones are formed by new osteons within the gap and by osteons present within the preexisting bone near the fragment ends. Cutting cones originating from both the gap and the fragment ends cross the fracture plane to unite the new lamellar bone within the gap to the fragment ends. With time, the new lamellar bone in the gap will become longitudinally oriented and reestablish the anatomic and mechanical integrity of the cortex.

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Direct healing (osteonal reconstruction) of the fracture can also occur in the presence of callus and is termed secondary osteonal reconstruction. With this method of bone healing, the fixation does not provide the stability necessary to reduce the deformation within the fracture gap to a level necessary for direct deposition of bone. Areas within the fracture plane are subject to instability and initially high levels of interfragmentary strain. Within these zones, bone resorption of the fragment ends occurs, which lengtlfens the fracture gap and reduces interfragmentary strain. Simultaneously external callus is formed and proceeds to stabilize the fragment ends. The fixation does provide the stability necessary for rapid bridging of the fracture line by external callus. Once the external callus unites, the deformation within the fracture gap is reduced to levels where bone tissue can survive between the fragment ends. If the fracture gap is less than approximately 1 mm wide, osteonal reconstruction of the cortex will then proceed as for gap healing described earlier. If the fracture gap is too wide, if the vascular supply is impaired, or if interfragmentary deformation does not allow for survival of osseous tissue, indirect bone healing will occur. Indirect bone healing (intermediate callus formation) is defined as that which occurs when osseous tissue is formed through the transformation of fibrous tissue or cartilage tissue.4 This method of bone healing occurs when interfragmentary deformation, impairment of blood supply, or width of the fracture gap will not allow direct formation of lamellar bone. Tissues are initially deposited within the fracture environment and subsequently prepare the fracture gap for survival of bone cells. Indirect bone healing has been divided into four stages: inflammation, soft callus, hard callus, and remodeling. 4 Inflammation begins immediately following fracture and persists until the initiation of fibrous tissue or cartilage formation. At the time of fracture, there is disruption of blood vessels with attendant hemorrhage and hematoma formation. The appearance of granulation tissue marks the beginning of a soft callus. Granulation tissue matures into fibrous tissue or fibrocartilage depending on local environmental conditions. At the periphery of the external callus, where blood supply is abundant, fibrous tissue is formed . Toward the center of the callus between the fragment ends, where blood supply is limited, fibrocartilage is formed. The amount of callus formed is directly dependent on the degree of motion present. The greater the degree of instability, the larger the diameter of the internal and external callus. The larger the diameter of the external callus (measured from the neutral axis of the bone), the greater is its ability to resist bending and rotation of the damaged bone. Soft callus tissues (fibrous tissue and fibrocartilage) do not have sufficient mechanical properties to decrease local deformation to a level conducive to osteoblast survival. To increase the stiffness of the fracture environment further, mineralization of the tissue begins and signifies the beginning of a hard callus. In fibrocartilage, mineralization of the matrix progresses from the fragment ends to the center of the fracture

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gap. In fibrous tissue, mineralization occurs between the collagen fibrils to become fiber bone. Mineralized fibrous tissue and fibrocartilage have the structural strength and stiffness to limit the gap deformation to acceptable levels whereby bone formation can begin. Soft callus undergoes endochondral ossification, in which osteoid is deposited on the scaffold of mineralized cartilage and mineralized fiber bone. The mineralized fibrous tissue and fibrocartilage are gradually replaced to form cancellous bone. The structural stiffness and strength of the cancellous callus is sufficient for a return to function once complete bridging has occurred. The healed bone, at this time, has a greater diameter than normal and may be misshapen. However, the cancellous structure is not permanent. Over the ensuing months or years, and through the stage of remodeling, cancellous bone is changed to longitudinally oriented lamellar bone and the bone contour is restored. The method of bone healing is determined by the stability of the fixator-bone composite and biological factors. The rigidity of fixation ultimately depends on the biomechanical characteristics of the fracture, the accuracy of reduction, and the amount of physiologic loading. Comparative experiments using a canine tibial osteotomy model have been performed to study the effects of mechanical environment on healing of fractures stabilized by external fixation. 12 When unilateral and bilateral fixators were compared, the more rigid bilateral frames resulted in osteotomy healing with less callus formation and more primary osteonal reconstruction across the osteotomy sites. 12 Another study, however, showed that static compression applied across an osteotomy, although more rigid, did not enhance healing compared with noncom pressed osteotomies in the same animals; periosteal new bone formed around both osteotomies. Some osteotomies of each group showed haversian remodeling through contact or gap healing. 12 Little is known about the influence of the timing of application of load to the healing process or of the effects of direction, magnitude, frequency, and application rate of load. Experimental work has shown that very short periods of appropriate dynamic mechanical stimulation can induce adaptive remodeling in intact bone. 32 Recent studies on the effect of controlled axial micromovement of external fixator stabilized fractures and osteotomies have shown that the fracture healing process is very sensitive to small periods of daily deformation applied axially within 2 weeks of fracture. 24 Although appropriate strain application may enhance healing, if boundaries of strain magnitude and force of application are exceeded, the healing process is inhibited. A recent clinical study of radial and tibial fractures repaired with external fixation showed that comminuted fractures healed by formation of periosteal and endosteal callus. 22 Fractures in which anatomic reduction was achieved with stable fixation showed radiographic signs of primary osteonal reconstruction. Open fractures tended to have more prolonged healing times, whereas fractures treated with closed reduction and stabilization tended to heal more rapidly.

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STRATEGIES TO DECREASE PATIENT MORBIDITY THROUGH APPLICATION OF BIOMECHANICAL AND BONE HEALING PRINCIPLES OF EXTERNAL SKELETAL FIXATION

Each orthopedic patient is different. In our clinic, we make a point of developing fracture fixation plans based on analysis of the biomechanical, biological, and clinical factors aHecting bone healing and return to normal locomotor function. Biomechanical factors that influence bone healing and return of normal locomotor function include number of limbs injured, patient size, and the ability to achieve load-sharing between the bone column and fixator. Weight bearing cannot be prevented when multiple limbs are injured; therefore implants are heavily loaded soon after implantation. Logically the incidence of complications often increases in such instances. Large patients subject their fixation to much greater loads and are predisposed to complications of premature pin loosening and fatigue failure of implants. Biological influences on bone healing include local and systemic factors. Local factors pertain to the area of the skeletal injury itself. Bone healing is more prolonged in open fractures than in closed fractures. Minimal soft tissue manipulation during fracture fixation is preferable to excessive manipulation (Le ., where possible, closed reduction is desirable). Fractures due to low-energy trauma such as spiral fractures with minimal soft tissue swelling have a better prognosis than high-energy comminuted fractures such as gunshot injuries. Systemic biological factors such as patient age and general health (hyperadrenocorticism, diabetes mellitus, hypothyroidism, etc.) obviously influence fracture healing and prognosis for return of normal locomotion. Prolonged healing, owing to local or systemic biological disease, predisposes patients to implant-associated complications owing, in part, to the increased number of loading cycles on the implant-bone composite before fracture union. Clinical factors must also be taken into account when developing a fracture fixation plan. Cooperative patients and compliant pet owners can ensure appropriate physical therapy will be performed without permitting undue load cycles to the fracture and implants. We have developed a group of strategies for maximizing pin-bone composite stiffness that we use when developing our fracture fixation plans for patients in our clinic. When all the biomechanical, biological, and clinical factors have been considered, a fixation and postoperative care plan that is appropriate for the individual patient is developed. In some fracture patients (when factors indicate prolonged healing, high implant loading, and poor patient compliance), each of the strategies is applied. In other patients (when factors indicate rapid healing, minimal implant loading, and good patient compliance are likely), fewer of the strategies are applied. Strategy 1. Surgeons should decrease pin stress (and interfacial tissue strain) by maximizing load-sharing between the fixator and the bone column. 1, 31

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Where possible (transverse fractures and osteotomies), achieving compression between main fracture segments is desirable. If compression is not possible, a fixator can be used in a neutralization fashion by accurate reduction and interfragmentary compression of butterfly fragments or oblique fractures with full cerclage wire or lag screws. In some instances, buttress fixation cannot be avoided. Such situations maximally "test" the structural design and material properties of the fixator-bone composite (pin-bone interface, fixation pins, pin clamps, and connecting rods) and require utmost attention to detail and decision making. Transverse fractures of the extremities (radius or tibia) can often be reduced in a closed fashion and compression achieved. In other instances, the mechanical advantages of increased load-sharing must be weighed against the biological disadvantages of open approaches to achieve anatomic reduction. As an example, a long oblique fracture of the tibial diaphysis may be stabilized by closed reduction and stabilization with a fixator; this represents a buttress situation, and there is no load-sharing by the bone column. In contrast, an open approach may be used and interfragmentary compression achieved with full cerclage wires and the fracture spanned with an external fixator; this represents neutralization fixation. The open surgical-approach affords load-sharing by the bone column through anatomic reconstruction of the fracture (neutralization), but at the expense of soft tissue trauma and introduction of contaminating bacteria into the surgical wound. In most instances, we find open reduction to warrant the biological insults only if significant mechanical advantages are gained over what can be accomplished with closed reduction and fixation. Strategy 2. Efforts should be made to maximize the number of fixation pins to four per major fragment. Maximizing the number of pins increases stiffness of fixator-bone composite, decreases the incidence of pin loosening, and decreases the cyclic stress applied to each fixation pin. lO, 12, 13 Although angled pins improve fixator-bone composite stiffness, surgeons must take into account that angled pins occupy more available bone stock, thereby limiting the overall number of pins that can be placed in each fracture fragment. Although addition of pins (up to four per fragment) is mechanically advantageous, each pin could represent a tract for introduction of contaminating bacteria to enter the wound and additional restriction on normal soft tissue movement over bone. In most instances, we prefer to maximize the number of fixation pins (up to four per fragment) while being careful not to penetrate muscle bellies during pin insertion. Not only is the incidence of premature pin loosening decreased when more pins are used, but if it occurs, we can often afford simply to remove the loose pin. Pins can be removed with sedation and minimal surgical field preparation. Conversely if very few pins are used initially and removing a loose pin is required, replacement of the pin with a new pin in a different implantation site is often necessary. Placement of new pins requires general anesthesia and thorough preparation of an aseptic surgical field.

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Strategy 3. Pin diameter should be maximized up to 20% the bone diameter. Maximizing pin diameter will increase pin stiffness and decrease interfacial tissue strain.l3, 31 The fixation clamps (pin-grippers) and the size of bone limit the size of pin that can be used. With Kirschner external fixation equipment, the small clamp accepts pins up to 2.4 mm (%2"), the medium clamp up to 3.2 mm (Vs"), and the large clamp up to 5.0 mm WI6"). Proportionate loss of structural strength occurs with each increase in circular cortical 'defect size greater than 20% the bone diameter. IS Strategy 4. Pins adjacent to the fracture should be placed approximately 1.5 to 2.0 cm from the fracture. Pins placed close to the fracture decrease the working length of the connecting rod, thereby increasing frame stiffness. Placement of pins too close to the fracture, however, may allow contaminating bacteria from the external environment to enter the area of the fractured bone and highly traumatized soft tissues. The remaining pins are customarily spread throughout the remaining bone stock of each fracture segment. Strategy 5. Connecting rods should be placed approximately a finger'S breadth from the skin.2 Minimizing the working length of each fixation pin decreases pin-bone interface stress and the likelihood of fatigue failure of fixation pins. 31 Some surgeons attempt to increase the bending stiffness of pins further by placing the pin clamps and connecting bars against the skin; significant patient morbidity results, however, when clamps impinging on the skin cause ulceration. Strategy 6. Frame stiffness should be maximized. Bilateral frames are stiffer than unilateral frames. 10, 16 Biplanar frames are stiffer than uniplanar frames. 16 Unilateral frames with two connecting bars are stiffer than if just one connecting bar is used. 16 As previously stated, increasing fixator-bone composite stiffness decreases the incidence of premature pin loosening and decreases the cyclic stress applied to each fixation pin.12 In most instances, we prefer to use a slightly stiffer frame configuration than we initially think may be necessary to achieve fracture union and restoration of locomotor function. Such an approach affords us the option of early destabilization of the frame if necessary (requires sedation only). Failure to use a stiff enough frame may lead to complications of premature loosening of pins, fatigue failure of pins, loss of fracture reduction, poor limb use, and delayed union and may require a second surgical procedure. Uniplanar frames (types Ia and II) are two to seven times more resistant to bending in the plane of the fixator than in the plane perpendicular to the fixator.s Unfortunately although uniplanar frames applied to the tibia are usually in a mediolateral plane, bending moments applied to the tibia during walking (in humans) are much greater in the craniocaudal plane. Resistance to bending in the plane perpendicular to the fixator can be improved with little additional bulk by minimizing the distance from the bone to the connecting rod, using two connecting rods, and maximizing separation of fixation pins within each fracture segment. S Alternatively addition of a unilateral frame in the craniocaudal plane to the unilateral frame in the mediolateral

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plane (making a unilateral-biplanar, type Ib) dramatically improves resistance to craniocaudal bending moments. 5 Because of the minimal soft tissue covering of the cranial surface of the tibia, the connecting rod can be positioned very close to the long axis of the bone, effectively reducing the working length of the fixation pins and increasing their stiffness. When using bilateral frames, use of primarily full-pins will maximize stiffness as compared with using some half-pins within the configuration. 10 If, however, we find that we have misdirected the placement of a full-pin such that we will not be able to connect it to the opposite fixation clamp, we will often leave it as a half-pin. In such instances, we will usually place another half-pin through the corresponding fixation clamp on the opposite connecting rod.

Strategy 7. Adding a second connecting rod to a unilateral-uniplanar frame will increase the compressive and bending stiffness of the fixator-bone composite.10, 16 This may be advantageous in such situations as gunshot fractures of the humerus or femur where the body wall does not permit bilateral frames to be used. The mechanical advantage of adding a second connecting rod must be weighed against the increased cost (fixation clamps are the most costly component) and increased bulk, but the ability to increase stiffness without adding extra pins and technical steps is attractive. Strategy 8. Do not use high-speed power drills for insertion of fixation

pins because thermal necrosis of bone causes premature fixation pin loosening. 17, 26 Insert pins by using a slow-speed power drill (150 rpm) or by predrilling holes with a twist drill bit (approximately 1 mm smaller than the pin diameter) followed by pin placement with a hand chuck. 17 Strategy 9. Use pins with sharp points. Dulled and dented pin points increase the temperature of bone during drilling and cause increased microstructural trauma to the bone. 26, 39 Strategy 10. Use threaded pins alone or in combination with non threaded pins. In general, non threaded pins are stiffer than threaded pins of equivalent outer diameter.6 On the other hand, non threaded pins have less holding power than threaded pins. 6, 29 Clinically combined use of threaded and nonthreaded pins within the same fixator provides structural stiffness while maintaining adequate holding power. 2, 3 Threaded and non threaded pin combinations reduce patient morbidity as compared with use of non threaded pins alone.3 There are recent reports of fatigue failure of Ellis pins (singlecortex, half-pin) during clinical use. 19, 28 In demanding cases involving multiple limb injuries, open fractures, highly comminuted fractures leading to buttress fixation, large-breed dogs, we avoid the use of Ellis pins. At present we prefer to use pins with a positive thread profile in such instances. Strategy 11. Appropriate postoperative management should be provided to minimize complications. Proper postoperative care is vital for consistent success with external fixation.

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SUMMARY

To decrease patient morbidity with external fixation, attention must be given to the type of fixator applied, the method by which it is applied, and the care provided for the patient after it has been applied. References 1. Aron ON: External skeletal fixation. Vet Med Rep 1:181-201, 1989 2. Aron ON, Toombs JP: Updated principles of external skeletal fixation . Comp Cont Ed Pract Vet 6:845-859, 1984 3. Aron ON, Toombs JP, Hollingsworth SC: Primary treatment of severe fractures by external skeletal fixation: Threaded pins compared with smooth pins. J Am Anim Hosp Assoc 22:659-670, 1986 4. Arnoczky SP, Wilson JW, Schwarz P: Fractures and fracture biology. In Slatter DH, ed: Textbook of Small Animal Surgery. Philadelphia, WB Saunders, 1985, pp 19391944 5. Behrens F, Johnson W: Unilateral external fixation. Methods to increase and reduce frame stiffness. Clin Orthop Rei Res 241:48-56, 1989 6. Bennett RA, Egger EL, Histand M, Ellis AB: Comparison of the strength and holding power of 4 pin designs for use with half pin (type 1) external skeletal fixation. Vet Surg 16:207-211, 1987 7. Black J: Deformation. In Orthopedic Biomaterials in Research and Practice. New York, Churchill Livingstone, 1988, pp 23-56 8. Black J: Mechanical properties. In Orthopedic Biomaterials in Research and Practice. New York, Churchill Livingstone, 1988, pp 57-81 9. Black J: Fixation. In Orthopedic Biomaterials in Research and Practice. New York, Churchill Livingstone, 1988, pp 267-283 10. Brinker WO, Verstraete MC, Soutas-Little RW: Stiffness stw;lies on various configurations and types of external fixators. J Am Anim Hosp Assoc 21:801-808, 1985 11 . Carter 0: SI: The international system of units. In Nordin M, Frankel VH, eds: Basic Biomechanics of the Musculoskeletal System, ed 2. Philadelphia, Lea & Febiger, 1989, pp xvii-xxiii 12. Chao EYS, Aro HT, Lewallen DG, Kelly PJ: The effect of rigidity on fracture healing in external fixation. Clin Orthop Rei Res 241 :24-35, 1989 13. Chao EYS, Pope M: The mechanical basis of external fixation. In Seglison 0, Pope M, eds: Concepts in External Fixation. Orlando, Grune & Stratton, 1982, pp 13-39 14. Churches AE, Tanner KE, Harris JD: The Oxford external fixator: Fixator stiffness and the effects of pin loosening. Eng Med 14:3-12, 1985 15. Edgerton BC, An KN, Morrey BF: Torsional strength reduction due to cortical defects in bone. J Orthop Res 8:851-855, 1990 16. Egger EL: Static strength evaluation of six external skeletal fixation configurations. Vet Surg 12:130-136, 1983 17. Egger EL, Histand MB, Blass CE, Powers BE: Effect of fixation pin insertion on the bone-pin interface. Vet Surg 15:246-252, 1986 18. Egkher E, Martinek H, Wielke B: How to increase the stability of external fixation units. Mechanical tests and theoretical studies. Arch Orthop Trauma Surg 96:35-43, 1980 19. Ellis AB, Dickason JM: Clinical trials of the Ellis partially threaded pin. Proc Vet Orthop Soc 14:2, 1987 20. Evans M: The Oxford external fixator. The 6th Annual Report of Oxford Orthopedic Engineering Centre, 1979 21. Gumbs JM, Brinker WO, DeCamp CE, et al: Comparison of acute and chronic pullout resistance of pins used with the external fixator (Kirschner splint). J Am Anim Hosp Assoc 24:231-234, 1988 22. Johnson AL, Kneller SK, Weigel RM: Radial and tibial fracture repair with external

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skeletal fixation: Effects of fracture type, reduction, and complications on healing. Vet Surg 18:367-372, 1989 23. Kasman RA, Chao EYS: Fatigue performance of external fixator pins. J Orthop Res 2:377-384, 1984 24. Kenwright J, Goodship AE: Controlled mechanical stimulation in the treatment of tibial fractures. Clin Orthop Rei Res 241 :36-47, 1989 25. Ling RSM: Observations on the fixator of implants to the bony skeleton. Clin Orthop Rei Res 21:80-96, 1986 26. Matthews LS, Green SA, Goldstein SA: The therIJIal effects of skeletal fixation-pin insertion in bone. J Bone Joint Surg [Am) 66:1077-1083, 1984 27. Nordin M, Frankel VH: Biomechanics of bone. In Nordin M, Frankel VH, eds: Basic Biomechanics of the Musculoskeletal System, ed 2. Philadelphia, Lea & Febiger, 1989, pp 3-29 28. Palmer RH, Aron DN: Ellis pin complications in seven dogs. Vet Surg 19:440-445, 1990 29. Palmer RH, Hulse DA, Polio F, et al: Pin loosening in external skeletal fixation: The effects of pin design and implantation site. Proc Vet Orthop Soc 18:50, 1991 30. Perren SM: Physical and biological aspects of fracture healing with special reference to internal fixation. Clin Orthop Rei Res 138:175-196, 1979 31. Pope MH, Evans M: Design considerations in external fixation. In Seglison D, Pope M, eds: Concepts in External Fixation. Orlando, Grune & Stratton, 1982, pp 109-135 32. Rubin CT, Lanyon LE: Regulation of bone formation by applied dynamic loads. J Bone Joint Surg [Am) 66:397-402, 1984 33. Schatzker J: Concepts of fracture stabilization. In Sumner-Smith G, ed: Bone in Clinical Orthopedics. Philadelphia, WB Saunders, 1982, pp 387-398 34. Schatzker J, Horne JG, Sumner-Smith G: The effect of movement on the holding power of screws in bone. Clin Orthop Rei Res 108:115-126, 1975 35. Schatzker J, Sanderson R, Murnaghan JP: The holding power of orthopedic screws in vivo. Clin Orthop Rei Res 108:115-126, 1975 36. Uhthoff HK: Mechanical factors influencing the holding power of screws in compact bone. J Bone Joint Surg [Br) 55:633-639, 1973 37. Uhthoff HK, Germain JP: The reversal of tissue differentiation around screws. Clin Orthop Rei Res 123:248-252, 1977 38. Vangness CT, Carter DR, Frankel VH: In Vitro evaluation of the loosening character-

istics of self tapped and nonself tapped cortical bone screws. Clin Orthop Rei Res 157:279-286, 1981 39. Wagenknecht M, Adrianne Y, DonkerwoIcke M, et al: Pin technology. In Coombs R,

Green SA, Sarmiento SC, eds: External Fixation and Functional Bracing. London, Orthotext, 1989, pp 143-147 40. Yanoff SR: Measurement of vertical ground reaction force in jumping dogs. Master's Thesis, Texas A&M University, 1991

Address reprint requests to Ross H. Palmer, DVM, MS Santa Cruz Veterinary Hospital 2585 Soquel Drive Santa Cruz, CA 95065