Etiology of Fractures

8/3/2019 Etiology of Fractures

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Etiology of Fractures Extrinsic CausesIntrinsic Causes Classification of Fractures by Type Incomplete Fractures

Complete FracturesClosed FractureOpen Fracture Classification of Fracture by Location Diagnosis of Fracture Pathology Associated with Fracture

A fracture is a dissolution of bony continuity with or without displacement of the fragments. It isalways accompanied by soft tissue damage of varying degrees, there are torn vessels, bruisedmuscles, lacerated periosteum, contused nerves. Sometimes there are injured internal organs andlacerated skin. The trauma to soft tissue must always be taken into consideration and is often

vitally more important than the fracture itself .(5)

Fractures have been classified by many authorities in the past. (1-4,6-8) This chapter attempts topresent a system of nomenclature appropriate for dogs and cats.

ETIOLOGY OF FRACTURESEXTRINSIC CAUSESDIRECT VIOLENCETrauma is the most common cause of fractures in small animals and is usually due to automobile

injury or falling from a height. Since direct trauma is rarely delivered in a calibrated amount to aspecific place, the resultant fracture is rarely predictable. The amount and direction of force willvary from accident to accident. Most fractures resulting from violent direct trauma are eithercomminuted or multiple.

INDIRECT VIOLENCEFractures due to indirect trauma are more predictable than those due to direct trauma. Generally aforce is transmitted to a bone in a specific fashion and at a "weak link" within the bone, causing afracture to occur.

BENDING FORCES

Bending fractures occur when force is applied to a specific focal point on a bone to the extentthat the traumatic force overcomes the elastic limit of the bone diaphysis. The initial effect of abending force is a cortical break opposite the site of the trauma. The periosteum will remainintact on the side of the force while tearing over the fracture on the opposite side. With additionalforce the entire bone snaps, with attendant tearing of vascular and soft tissue structures within oron the diaphysis. Bending fractures are generally oblique or transverse, or they may have abutterfly fragment. (Example: A dog running across a field steps into a gopher hole with the hindlimb; the edge of the hole is a fulcrum producing a bending fracture of the midshaft tibia.)

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TORSIONAL FORCESTorsional fractures occur when a twisting force is applied to the long axis of a bone. Usually thisis a result of one end of a bone being placed in a fixed position while the other end of the bone isforced to rotate. The resulting fracture will be a very long spiral with sharp points and oftensharp edges. It is possible for the sharp points or edges to compromise soft tissues or to cut

through the skin and result in an open fracture. Torsional forces generally result in short or longspiral fractures. (Example: A cat jumping from a garage roof to a fence misjudges the distanceand catches its hock in the fence. The resulting force of its body twisting against the fixed lowerextremity results in a spiral fracture of the tibial diaphysis.)

COMPRESSION FORCESCompressive forces along the long axis of a bone may force the smaller diaphyseal ormetaphyseal portion of a bone to impact into the larger epiphysis: bony substance is therebycrushed. Similarly a compressive force directed along the axis of the spine may result in collapseof a vertebral body. For compressive force to result in fracture, one end of a bone must be in afixed position while the other end is forced toward the fixed end. Compressive forces result in

impacted fractures or compression fractures. (Example: A large breed puppy jumps for a frisbeeand in landing forces the hock plantigrade into the ground. The full weight of the dog thencrushes the proximal tibial epiphysis over the proximal tibial metaphysis.)

SHEARING FORCESA shearing fracture is caused by a force transmitted along the axis of a bone, which is thentransferred to a portion of the same bone that lies peripheral to the axis or across a joint to otherbones that are not protected by the axis of the bone. The force shears off that bony portion unableto continue transmission of the force along the axis. The fracture line in a shear fracture will beparallel to the direction of the applied force. Shearing forces result in the fracture of bonyprominences not placed along the direct axis of a diaphysis. (Example: An immature miniaturebreed dog is dropped from its owner's arms to a hard surface. The force transmitted up the radiusand ulna, across the elbow joint and into the distal humerus will shear off the lateral humeralcondyle.)

INTRINSIC CAUSESFRACTURES DUE TO MUSCULAR ACTIONFractures caused by violent contraction of a muscle are called avulsion fractures. They mayoccur because of violent isometric contraction but are associated more commonly with traumathat results in forceful muscular shortening. These fractures frequently occur in immatureanimals while the physeal plate remains open. Such muscular forces are more likely to separate acartilaginous union than the eventual bony union of mature animals.

Avulsion fractures affecting bony prominences that serve as the major origin or insertion of amuscle are seen routinely. The processes commonly avulsed include the acromion, scapulartuberosity, greater humeral tubercle, olecranon, ischial tuberosity, greater trochanter, tibialtuberosity, and the calcaneus of the fibular tarsal bone.



PATHOLOGIC FRACTURESPathologic fractures occur because of underlying bony or systemic disease that causes one,many, or all bones of an animal's skeletal system to be abnormal and thus more susceptible tofracture. Pathologic fractures may occur from any type of trauma: bending force, torsional force,compressive force, or shearing force. Often the only force necessary to cause fracture is the

animal's weight; thus, spontaneous fracture occurs without overt trauma.

Pathologic fracture may occur through any of the following types of bony pathology: neoplasia,bone cysts, osteoporotic bone caused by secondary NHPO, nutritional hyperparathyroidism,localized bone infection (osteomyelitis), osteoporotic bone caused by disuse following prolongedexternal fixation or removal of a rigid internal device(Fig. 11-1).

A pathologic fracture can occur in any bone, in any location within a bone, and take any shape.The diagnosis of underlying pathology is usually of more importance than immediate bonefixation. Once the pathologic basis for the fracture has been diagnosed and specific correctivemeasures initiated, the fracture or fractures can be treated. Treatment of all pathologic fractures,

including those due to neoplasms, can be successful.

FIG. 11-1 Pathologic fracture. Fibrosarcoma of the distal femoralmetaphysis in a dog.

CLASSIFICATION OF FRACTURES BY TYPEFractures are classified into many types based on the severity of the fracture, whether itcommunicates through the skin, the shape of the fracture line, or the anatomical location of thefracture within an individual bone. All systems are compatible and of necessity overlap.

INCOMPLETE FRACTURESAn incomplete fracture implies that a bone has not completely lost continuity; some portion of the bone remains intact. There are several types of incomplete fractures.

GREENSTICK FRACTUREAs the name implies, a greenstick fracture resembles the break that results when a supple greenbranch of a tree is bent and breaks incompletely. Usually the side opposite the bending forcefractures completely, while the side under the force remains intact. In the immature animal with

similarly supple elastic bone, a bending force will produce the incomplete fracture. Since aportion of the bone cortex remains intact, this fracture cannot override and result in limbshortening; however, the limb may deform along its axis at the point of the bending force (Fig.11-2).

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FIG. 11-2 Incomplete fracture of the femoral diaphysis.

FISSURE FRACTURECracks or fissure lines will occur when direct trauma is applied to any long or flat bone.Generally the fissures are formed in one cortex of the bone and are covered by an intactperiosteum. Bones may have single or multiple fissure lines of any configuration: transverse,oblique, spiral, longitudinal, or radiating from a central point. Since fissure fractures occur onlyin a single cortex and represent an incomplete fracture, the fractured bone should maintain itsnormal shape.

DEPRESSION FRACTUREDepression fractures represent areas in which multiple fissure fracture lines intersect. Withsufficient force, the entire area will depress from the direction of force. This usually occurs in thecalvarium, the maxilla, or the frontal bone areas of the head.

COMPLETE FRACTURESComplete fractures are indicated by the complete loss of bony continuity, allowing overridingand deformation. Complete fractures are far more common than incomplete fractures. They maybe classified further by the shape of the fracture line. The following system describes completefractures.

TRANSVERSE FRACTURETranverse fracture implies a fracture line that is transverse to the long axis of the bone.Transverse fractures may be relatively smooth or may be rough or have deep teeth on thefractured surfaces. Most are caused by bending forces. Roughness simplifies anatomicalalignment and increases the likelihood of rotational stability once reduced. Once these fracturefragments have been reduced, fragment override should not occur (Fig. 11-3).

OBLIQUE FRACTUREOblique fracture implies a fracture line that is oblique to the long axis of the bone. The twocortices of each fragment are in the same plane without spiraling. The edges of an obliquefracture may be rough but are usually smooth. The cortical edges are flat, rather than sharp.These fractures generally result from bending, with superimposed axial compression. As a result

of the obliquity of the fracture line, this fracture tends to override or rotate unless traction ismaintained throughout the period of healing (Fig. 11-4).

SPIRAL FRACTURESpiral fracture indicates a fracture line that spirals along the long axis of the bone; it is caused bytorsional twisting or rotational forces. Spiral fractures tend to have extremely sharp points andedges, which frequently accompany soft tissue trauma osr an open fracture. Reduction of spiral




































fractures is difficult without constant traction or internal fixation, since these fractures tend tooverride and rotate into deformity (Fig. 11-5).

COMMINUTED FRACTUREComminuted fracture implies at least three fracture fragments, the fracture lines of which

interconnect. The individual fracture lines that form the comminuted fracture may be transverse,oblique, or spiral. Comminuted fractures are generally caused by high-energy trauma, as typifiedby automobile accidents, and are a common type of animal fracture (Fig. 11-6). Comminutedfractures are difficult to reduce and fix because they have no inherent stability. Constant externaltraction and alignment or internal fixation is required.

FIG. 11-3 Transverse fracture line. Drawing represents a reduced transversefracture of the midshaft femoral diaphysis.

FIG. 11-4 Oblique fracture line. Drawing represents a reduced obliquefracture of the midshaft femoral diaphysis.

FIG. 11-5 Spiral fracture line. Drawing represents a reduced spiral fractureof the midshaft femoral diaphysis.

FIG. 11-6 Comminuted fracture lines. Drawing represents a reducedcomminuted fracture of the midshaft femoral diaphysis.

FIG. 11-7 Multiple fractures. Drawing represents a reduced femoral neckfracture and a reduced transverse fracture of the distal femoral metaphysis.

MULTIPLE FRACTUREMultiple fracture implies three or more fracture fragments in a single bone; however, unlikecomminuted fractures, the fracture lines do not interconnect. The individual fracture lines may beof any shape. Typically this term describes two completely independent fractures affecting thesame bone, such as an oblique fracture of the proximal femur and an epiphyseal fracture of the




































distal femur. Neither of these fractures interconnects. Reduction and fixation of a multiplefracture requires two separate reductions and fixations (Fig. 11-7).

IMPACTION FRACTUREDistinguishing between impaction fracture and compression fracture is difficult; however,

because both terms are used routinely in orthopaedic texts, the difference will be clarified. Animpacted fracture implies a fracture in which a bony fragment, generally cortical, is forced orimpacted into cancellous bone. Typically this occurs at the ends of long bones. Reduction of such fractures requires traction to disengage the fragments and fixation to hold the fragmentsapart. If, after fracture, malalignment is untreated, bone shortening will occur because one endhas impacted into the other. This is an uncommon fracture in small animals.

COMPRESSION FRACTIONCompression fractures are similar to impaction fractures, but the term is used to describe afracture in which cancellous bone collapses and compresses upon itself. Typically this occurs invertebral bodies following trauma to the spine. Compression fractures are rarely reduced, since

the bone within the fracture area has been destroyed by the crushing. These fractures are stableand heal in place; however, shortening occurs as a result of compression (Fig. 11-8).

FIG. 11-8 Compression fracture. Drawing representsan unreduced compression fraction of a lumbarvertebral body.

CLOSED FRACTUREA closed fracture implies a fracture that remains encased within the skin and musculature thatsurround it. No wound or mucosal membrane overlies the fracture. The fracture does notcommunicate with the outside environment. Most fractures in animals are closed. A synonymfound in older literature is "simple fracture" (Fig. 11-9, A)

OPEN FRACTUREUnlike a closed fracture, the open fracture communicates with the outside environment. Thismay occur through a large wound in the soft tissue and skin or through a tiny puncture wound.Regardless of wound size, any fracture that has communicated with the outside is considered anopen fracture. Of greatest significance is the potential for contamination of the fracture itself (Fig. 11-9, B). A synonym found in older literature is "compound fracture."

CLASSIFICATION OF FRACTION BY LOCATIONFractures may be classified by their anatomical location in relation to a specific bone. Identifyinga fracture by location does not indicate whether the fracture is open or closed, nor does it indicatethe type of fracture: transverse, oblique, spiral, or the like. The systems of classifying by typeand classifying by location are compatible and should be used together. FIG. 11-8 Compressionfracture. Drawing represents an unreduced compression fracture of a lumbar vertebral body.



















FIG. 11-9 (A) Closed reduced oblique fracture of the midshafttibial diaphysis. (B) Open unreduced oblique fracture of themidshaft tibial diaphysis.

DIAPHYSEAL FRACTUREFor purposes of description, fractures are termed midshaft if they occur near the axial center of the diaphysis. All other fractures of the diaphysis are referred to by breaking the diaphysis intoequal thirds. Therefore, fractures can be proximal third, middle third, or distal third of thediaphysis. A proper description would be closed, transverse fracture of the proximal thirddiaphysis of the femur. This classification should suggest a fracture within the skin, as well asthe shape, anatomical location, and the bone fractured.

METAPHYSEAL FRACTUREAny fracture within the anatomical metaphysis of a long bone is referred to as a metaphysealfracture. For a clearer description the terms proximal or distal should be added, such as a closed,oblique fracture of the distal femoral metaphysis. Since most metaphyseal fractures are throughcancellous bone, they generally heal rapidly.

FRACTURE OF THE EPIPHYSEAL PLATEFracture of the epiphyseal plate occurs in immature animals during the time that the epiphysealplate remains open and cartilaginous. Fracture occurs through the zone of hypertrophied cartilagecells. Referral to such fractures should specify the proximal or distal epiphyseal plate. In matureanimals, such fractures are called physeal fractures or fracture of the physis. Fractures of theepiphyseal plate are classified further to accurately describe their shape and severity of thefracture. The method of Salter- Harris is the standard classification for all species.(7) (See Figs.34-1 through 34-6.)

Type I-Epiphyseal separation: there is displacement of the epiphysis from the metaphysis at thegrowth plate.Type II-A small corner of metaphyseal bone fractures and displaces, with the epiphysis displacedfrom the metaphysis at the growth plate.Type III-Fracture is through the epiphysis and part of the growth plate, but the metaphysis isunaffected.Type IV-Fracture is through the epiphysis, growth plate, and metaphysis. Several fracture linesmay be seen.Type V-Impaction of the epiphyseal plate occurs, with the metaphysis driven into the epiphysis.

With each progressive type, the fracture described becomes increasingly difficult to treat andcarries a poorer prognosis for return to normal function.

EPIPHYSEAL FRACTUREIn the mature animal with closed growth plates, fractures of the epiphysis are termed epiphysealfractures. They should be classified further by describing them as fractures of the proximal ordistal epiphysis.









CONDYLAR FRACTURECondylar fractures occur in mature animals and affect the distal ends of the humerus or femur, orthe proximal tibia. Since anatomically a condyle is composed of metaphysis, physis, andepiphysis, this descriptive classification system is used instead of the previous three. Condylarfractures are further defined as medial or lateral, depending on the aspect fractured. If both

condyles fracture off the shaft as a unit, the fracture is termed supracondylar. Both condyles mayfracture from the shaft and from each other. This is a supracondylar/intercondylar fracture andmay be classified as a "V," "Y," or "T" fracture to better describe the shape of the fracture lines(Fig. 11-10). Any fracture of a condyle reflects potential problems if fracture of the joint surfacehas occurred.

ARTICULAR FRACTUREArticular fracture indicates that the subchondral bone and articular cartilage are involved in afracture. Such a fracture may be classified further by indicating which bone (proximal or distal)or which specific joint is fractured. Intra-articular fracture of the knee is nonspecific; descriptionmust specifically indicate fracture of the femoral or tibial component. Articular fracture is

synonymous with intra-articular fracture and means fracture within a joint. The term periarticularfracture is used to refer to fracture close to, but not into, the joint. The term could be replaced byepiphyseal fracture. Articular fracture requires perfect anatomical reduction and fixation toprevent secondary degenerative joint disease.

AVULSION FRACTUREAvulsion refers to a fracture of intrinsic etiology, generally caused by muscular contraction. Theprominences that fracture are usually separate centers of bone formation referred to asapophyses. Avulsion fractures are classified by the prominence that has been avulsed, such asavulsion of the greater trochanter. Avulsion fractures tend to displace in the direction of themuscle pull that caused the fracture. Reduction and fixation is difficult and requires constanttraction or internal fixation.

FRACTURE-DISLOCATIONFracture-dislocation describes joint fractures that produce joint instability sufficient to result insimultaneous subluxation or luxation of the affected joint. This classification is incomplete, sincefracture-dislocation of the shoulder indicates dislocation of the shoulder but does not indicatewhich bone, the scapula or the humerus, is fractured. Therefore, a more descriptive classificationof the fracture must be given. Fracture-dislocations can be difficult to treat because theyrepresent intra-articular fracture plus supporting tissue laxity. When fracture and dislocation arefound together, the prognosis is poorer than if each problem occurred separately.

DIAGNOSIS OF FRACTUREIn most instances the clinical signs associated with fracture make diagnosis uncomplicated.Although the owner of an animal often will have observed the fractured bone, locating a fracturecan at times be difficult. In these instances, the practitioner needs a systematic, logical approachto diagnose the fracture.






FIG. 11-10 Condylar fractures. (A)Lateral humeral condyle fracture. (B)Intercondylar and supracondylar fracturesof the distal humerus (a "T" fracture). (c)Intercondylar and supracondylar fractures

of the distal humerus (a "Y~ fracture).

DYSFUNCTIONDysfunction is most commonly exemplified by lameness. In the orthopaedic examination thefocal site of the lameness must be found and the diagnosis pursued. Dysfunction may alsoinclude paralysis with spinal fracture, unconsciousness accompanied by cranial fracture, ormasticatory dysfunction with mandibular fracture.

Impairment or loss of function is a constant sign of complete fracture and is the result of pain orloss of mechanical support. Only in cases of incomplete or impacted fracture may some weightbe borne by the bone. The careful observer will determine the difference between loss of

function due to pain alone and that due to inability to bear weight. The smaller the animal, themore difficult it is to make this distinction. Toy Pekingese, Pomeranians, and all cats requireconsiderable care in determining the presence of this sign.

PAINPain over the site of fracture is common. In incomplete fractures this may be the only clinicalindication. Direct tenderness can be misleading, since it may be due to a contusion or other softissue damage caused by a blow. Indirect tenderness is a more accurate sign of fracture. It isproduced by pressure in the long axis of the bone exerted at its two extremities. If there is a breakin the continuity of the shaft, such pressure will cause pain at the fracture site that is quitedistinct from the pain of injured soft tissue parts. If an animal is examined during the state of

local tissue shock, that is, within 20 to 30 minutes after the accident, pain may not be aconspicuous sign.

LOCAL TRAUMAExamination of the area around a fracture may demonstrate swelling, hematoma, contusion, orlaceration if the fracture is open. Often because of extreme swelling, the examiner will be unableto palpate crepitation. Local swelling, although present in many other conditions, is one of themost constant signs of a fracture. Immediately after injury the swelling may be sharply outlinedas a result of bleeding from the bone and the soft parts. An indistinctly outlined swelling thatoccurs later is caused by edematous infiltration. Generally the swelling increases for 24 to 48hours, then gradually subsides (particularly under treatment). When applying bandages and

splints immediately following fracture, it is important to bear in mind that swelling will subside.

ABNORMAL POSTURE OR LIMB POSITIONINGAbnormalities of positioning, when of acute onset associated with trauma, usually reflect afracture. Deformity, a deviation from the normal anatomical structure, may be caused bydisplacement of the bony framework as in a fracture or dislocation, but it may also be caused bychanges in configuration due to a neoplasm. The displacement of bone fragments that producesdeformity in a fracture may be angular, longitudinal, or rotational. Longitudinal displacements




may cause shortening, referred to as overriding, or may result in separation of the fragments,termed distraction (e.g., fractures of the olecranon). In most cases the primary displacement isdetermined by the direction and force of an injury and is maintained and often increased by thecontraction of muscles. If in doubt about positioning, comparison with the opposite limb or sideof the body part is advised.

CREPITUSCrepitus is a sign of fracture that is considered pathognomonic. Bony crepitus is the grittingsensation transmitted to the palpating fingers by the contact of the broken bone ends on eachother. There are other forms of crepitus (pseudocrepitus) such as occurs in some cases of arthritis, partial luxations of the patella, or luxations of the coxofemoral joint. The absence of crepitus does not necessarily indicate the absence of a fracture. The interposition of a piece of soft tissue between the fragments will prevent crepitus. It is also absent when the ends of thebones are so far apart that they cannot be brought into contact, or when they are impacted.Crepitation should be elicited with the utmost precaution because of the danger of causingfurther damage to bony fragments and surrounding soft tissue. Vigorous palpation, which may

turn a routine closed fracture into a contaminated open one should be avoided.ABNORMAL MOBILITYA false point of motion is also pathognomonic. It occurs if there is a complete fracture of theshaft of a long bone; it does not occur in an incomplete or impacted fracture. Mobility near a joint may be difficult to differentiate from normal or abnormal mobility of the joint itself. Inorder to avoid additional trauma, the same precaution should be taken in eliciting this symptomas in eliciting crepitus.

RADIOGRAPHIC SIGNSFracture, either diagnosed or suspected, should be documented by radiography. At least two

views including the joints above and below the fracture are needed. Fracture of joints or specialanatomical locations may require additional radiographs or special positioning. Radiographsshould be read on a well-illuminated flat surface. If questions about anatomical structures exist,the opposite limb or side of the body may be radiographed for comparison. The specificradiographic signs of fracture include those listed below: A break in the continuity of a bone Aline of radiolucency when the fragments are distracted A line of radiopacity when the fragmentsare compressed or superimposed (Fig. 11-11)

OTHER SIGNSAlthough all of the above signs do not always occur in all fractures, combinations of these signsare always present. As time elapses between the time of the trauma and the time of treatment,symptoms change in accordance with the changes at the fracture site. Miscellaneous signsassociated with fracture include the following:

Fever. Elevated temperatures are seen routinely 24 to 48 hours following a fracture and reflectthe response to breakdown of the hematoma.







Anemia. Medullary arteries are high-pressure vessels, and significant hemorrhage can occur withfracture. Large dogs may lose 200 ml to 300 ml of blood into the hematoma. Animals withmultiple bone fractures can lose this amount of blood into each hematoma.

Shock. Hypovolemic shock can readily occur with severe fracture or concomitant vascular

lacerations. Shock may lead to death following severe blood loss into a fracture site.

Nerve injury. Depending on the location of the fracture or its severity, peripheral nerves can beinvolved.

Necrosis or gangrene. In instances of fracture and simultaneous vascular laceration or occlusion,necrosis of distal extremities may occur. This usually occurs several days following fracture.

Fat in synovial fluid. This sign may indicate presence of an articular fracture; however, anytrauma to a joint may result in fat in the synovial fluid. If fat is found and the animal remainslame, further studies may be needed to pursue the diagnosis of fracture.

FIG. 11-11 Radiographic signs of fracture. Note both a line of radiolucency where fragments are distracted and two lines of radiopacitywhere fragments are superimposed.

PATHOLOGY ASSOCIATED WITH FRACTUREMany specific fracture types have soft tissue injuries associated with them. Pelvic fracture mayresult in laceration of the bladder, prostate, pelvic urethra, or major vessels and nerves. Fracturedribs routinely accompany hemothorax, pneumothorax, or laceration of the lung parenchyma.Fracture of the axial skeleton can be expected to compromise the brain, brain stem, or spinalcord.

There may also be associated injury to the surrounding soft tissue produced by the trauma thatcaused the fracture. It is important to remember that skin, muscles, periosteum, tendons, nerves,and vessels over the fracture absorbed the same force as the fractured bone. Any or all of thesestructures may be severely damaged at the time of impact.

Trauma sufficient to cause fracture may also produce whole body manifestations. Whileautomobile trauma can cause a fractured femur, the entire animal is involved and the likelihoodof shock is great. The brain or spinal cord may have contused within its bony case and becomeedematous. Fat embolization from the fracture site may occur and produce respiratory difficulty.Hemorrhage at the fracture site may be minimal, but a ruptured abdominal organ may result inblood loss sufficient to cause death.

In summary, when examining an animal with a fractured femur, it is important to remember thatthe entire animal may need treatment, as well as the fracture. Every fracture is part of a




functioning animal, and although the fracture may be obvious it is only a small portion of theproblems present because of the trauma sustained.

Bone as a Material Bone Function Versus Bone Failure

Mechanical Concepts Fracture Etiology TensionCompressionShearBendingTorsionFailure Under Combined LoadingEnergy to Failure Biomechanics of Fracture Reduction Biomechanics of Fracture Fixation

External ImmobilizationExternal Skeletal FixationInternal FixationIntramedullary FixationMathematical ModelingPlate-Induced OsteopeniaBiomechanics of Surgical Screws Biomechanics of Fracture Healing The Role of Functional Weight BearingInternal FixationComparative Biomechanical Studies: Healing Relative to Treatment Modality

Biomechanics is defined as mechanics applied to biology. Mechanics, in turn, is the analysis of any dynamic system, be it the relative motion of quanta and subatomic particles or the motion of galaxies. The term mechanics was used as early as 1638 by Galileo in a treatise describing force,motion, and the strength of materials. (1,2,38,89) Galileo and his colleagues Harvey, Santorio,and Descartes were unknowingly the early pioneers in biomechanics, basing their biologicdiscoveries on physical principles, astute observation, and quantitative analysis. Briefly, theirmost notable contributions include the discovery of blood circulation, development of themicroscope, theoretic mathematical modeling of animal structure, and early studies onmetabolism. Since its inception, the scope of biomechanics has expanded immensely, nowincorporating all of continuum mechanics in physiology and the mechanics of medical

application in the health sciences: from clinical problems in the cardiovascular system toquantitative physiology and rheology of biologic tissues to orthopaedic implants and thekinematics of the musculoskeletal system. It is the purpose of this chapter to focus on some of the more important biomechanical principles relating to fracture etiology, reduction, and repair.

BONE AS A MATERIALBefore discussing the mechanical properties of bone and the influence of forces thereon, it is











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important to have a solid understanding of the anatomy of bone. This topic has been coveredelsewhere in this book (Chapter 1) and should be reviewed. Compact (or cortical) bone is a"composite" material in the true sense of the word. Two thirds of the weight (and approximatelyone half of the volume) of compact bone consists of an inorganic material having a compositionsimilar to the formula for hydroxyapatite, namely, 3Ca3 (PO4)2Ca (OH)2. The remainder is

composed of organic material, mainly collagen. Collagen molecules band together in an orderlysequence to form collagen fibrils, which in turn run parallel to each other to form collagen fibers.Inorganic hydroxyapatite crystals (approximately 200 nm long and 50 nm x 50 nm in crosssection) attach at specific sites along the collagen fibril.

Collagen fibers are found to have a characteristic orientation relative to bone type. That is, inwoven fiber bone, collagen fibers are arranged in a randomly tangled array, whereas in lamellarbone collagen fibers run parallel within any given lamella. The parallel collagen fibers insuccessive lamellae are oriented approximately at right angles to each other (see Fig. 1-16). Aconcentric, cylindrical arrangement of lamellae constitutes the haversian system or osteon (seeFig. 1-16). Circumferential lamellae form the bone adjacent and parallel to the bone surface,

while interstitial lamellae represent remnants of old haversian systems with typical osteonalcharacteristics.

Virtually all bone is microscopically lamellar in nature. However, the degree of porosity yieldstwo macroscopically distinct types. Figure 12-1 shows a light photomicrograph of cortical bonein transverse section and, on the right, an electron photomicrograph of cancellous bone.(35) Themajor difference between the two is the relative porosity, which for cortical bone varies from 5%to 30%, whereas that for cancellous bone is 30% to over 90%.

In general, bone is a good composite material, having a strength higher than either of itscomponents, apatite or collagen. The softer (low-modulus) collagen prevents the stiff (high-

modulus) apatite from undergoing brittle fracture, while apatite acts as a rigid scaffold to preventcollagen from yielding. Not surprisingly, the mechanical properties of bone are as complex andvaried as the anatomy and composition. Seemingly simple properties such as bone strength,stiffness, and energy absorption to failure depend not only on material properties of bone (e.g.,inherent composition, microscopic morphology of bone components, bonds between fibers andmatrix and bonds at points of contact of fibers) but also on structural properties (e.g., geometryof whole bone, bone length, and bone curvature). Furthermore, it is well known that the materialstrength of bone varies with the age, sex, and species of animal under investigation and with thelocation of bone, such as femur versus humerus. In attempting to assess structural and materialproperties of bone using mechanical testing techniques, additional variation in bone strength mayresult from factors such as the orientation of load applied to the bone (since bone is anisotropic),strain rate (rate of deformation), and testing conditions, including tension versus compression,bending versus torsion, wet bone versus dry bone. Mechanical testing may yield even widerdispersion in results when specifically applied to the material and structural properties of healingbone. Amount of callus, type of callus, and degree of callus reorganization all affect themechanical assessment of bone healing.

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Clearly, these complexities (compositional, anatomical, mechanical, and experimental)necessitate careful scrutiny on the part of the orthopaedist in extracting clinical relevance fromthe available literature on bone and bone-healing biomechanics.

FIG. 12-1 (A) Reflected light photomicrograph of cortical

bone from a human tibia. (original magnification x 40)(B) Scanning electron photomicrograph of cancellousbone from a human tibia demonstrates marked porosityrelative to cortical bone. original magnification x 30)(Frankel VH, Nordin M Basic Biomechanics of theskeletal System. Philadelphia Lea & Febiger, 1980)

BONE FUNCTION VERSUS BONE FAILUREBones are linked together in an orderly array to form the skeletal system and in this regard havethree primary biomechanical functions: to encase or partially surround vital internal organs, thus

affording protection; to act as rigid kinematic links and attachment sites for muscles (intrinsicforce generators); and to facilitate body movement by means of well-lubricated (low-friction) joints. Given the importance of these functions to survival, bone, through evolution, hasdeveloped two unique properties: bone remodels to meet functional need, so-called Wolff'slaw(100) (1884); and bone has the capacity for repair. It is the biomechanical optimization of thislatter property that has most relevance to the practice of veterinary orthopaedics and thereforewill receive further elaboration in the sections that follow.

Bone, in performing its function, must withstand a complex pattern of imposed forces. In a staticsituation, bone acts largely to resist the forces of gravity, supporting the weight of the body andthe attendant muscular activity necessary to maintain a given static posture. In a dynamic mode,

however, such as during locomotion or athletic activity, these forces may be magnified manyfold and may be omnidirectional.

Bone, in function, experiences two types of imposed forces. In general, intrinsic forces may beconsidered physiologic and are imparted to bone through articular surfaces by means of ligaments surrounding joints and at tendinous sites of muscle insertion. Under normalcircumstances such forces sustain ground reaction forces during posture and gait and only underunusual circumstances do they approach the inherent breaking strength of bone. Extrinsic forces,on the other hand, originate from the environment and, unlike the intrinsic system, have nolimitation on magnitude or direction of application, for example, automobile impact. Clearly it isthese nonphysiologic forces that have the greatest potential to result in catastrophic bone failure

(fracture) and that must be understood to evaluate the biomechanics of fracture etiology.

Intrinsic and extrinsic forces act to cause microscopic deformations of bone. The degree of deformation is dependent on the magnitude of the imposed force, the geometry of the bone (size,shape, diameter, curvature), and the material properties of bone (cortical versus cancellous). It isintuitive that should the magnitude of imposed forces on bone exceed the ultimate strength of that bone, catastrophic failure will result. More specifically, bone fracture occurs at the point at




which the energy-absorbing capacity of bone is exceeded (KE = I/2 mv2). This point will bediscussed further in the section to follow.

MECHANICAL CONCEPTSThe biomechanics of bone in its normal and healthy state are integral to an understanding of fracture etiology and to the development of optimum repair modalities. The purpose of thefollowing section is to introduce some mechanical concepts that were developed for evaluationof inanimate materials systems but that may appropriately be applied to bone. Hopefully thisinformation will engender a biomechanical vocabulary common to both engineers and surgeons.

As stated previously, bone is a solid material and as such experiences forces and resultantdeformations in performing its function. In this context, force and load are used synonymously todefine the magnitude of the vector quantity (force) that acts to deform the structure, for example,the whole bone. In contrast, stress refers to the distribution of applied force over the cross-

sectional area of whole bone specifically.STRESS=FORCE/AREA

The relationship of force to stress will receive further attention shortly. However, it should berecognized that the two parameters (force and stress), although related, are not synonymous.

The measure of a material's deformation (e.g., elongation) can be expressed either as an absolutechange in length (delta L) or as normal strain, the change in length per unit initial length (deltaL/Li).

In contrast to normal strain, shear strain is the measure of the amount of angular deformation,alpha, in response to the application of shear force. Stress is usually expressed in units of newtons per centimeter squared (N/cm2), newtons per meter squared (N/m2), pascals (1 Pa= 1N/m2) or megapascals (1 mega Pa= 1 x 10^6 Pa). Normal strain is a dimensionless unit; forexample, centimeters/centimeter, and is expressed as a percentage while shear strain is expressedin radians (1 radian is approximately 57.3¡).

Stress and strain having been defined, it is possible to introduce the two most importantmechanical properties of bone as a material, namely, strength and stiffness. First, however, it isimportant to recognize the distinction between structural parameters and material parameters. Forexample, consider the situation depicted in Figure 12-2 in which a tibia and its correspondingfibula are loaded independently to failure in tension. It is obvious that the fibula will exhibitfailure at a lower load (force) than the tibia because of its reduced cross-sectional dimension.Under such a circumstance the tibia is said to have a higher structural strength than the fibula. If,however, the applied loads at failure are normalized relative to cross-sectional area of bone (F/A= stress) and if instead of absolute deformation (delta L), the strain (delta L/ Li) is measured, thematerial strength of the tibia and fibula can be shown to be comparable. This distinction isgraphically illustrated by examining a plot of the pertinent data on a "load-deformation" curverelative to a "stress-strain" curve (Fig. 12-3). It is apparent from the idealized stress-strain curve











that per unit bone volume, the tibia and fibula have similar material strength in spite of theirobvious structural differences. Such concepts must be appreciated when making comparativeassessments of bone strength, such as dog bone versus human bone, cortical bone versuscancellous bone, or cortical bone versus fracture callus.

FIG. 12-2 Tibia and fibula loaded to failure independently in tension toevaluate mechanical properties. It is intuitive that the fibula will fractureat a lower load than the tibia.

Strength, then, as a material parameter is defined as the ultimate stress at which failure occurs,whereas strength defined structurally is the ultimate load (or force) at which failure of the systemoccurs. This distinction becomes extremely important when discussing fracture healing. Forexample, on a materials basis (per unit volume) fracture callus has a strength far inferior tonormal bone yet on a structural basis, because of its abundant volume (cross section), may havestructural strength approaching that of whole intact bone.

To adequately define stiffness as a mechanical property, the stress-strain curve requires furtherelaboration. Figure 12-4 is an idealized stress-strain curve of a machined specimen of a ductilemetal having known cross section loaded in tension. As the specimen is loaded in tension frompoint A, its stress strain behavior proceeds in a linear fashion to point B, the yield point. If theload were removed at point B, the metal would return to its original undeformed length alongpath B-A; this is termed "elastic" behavior. The slope of the stress-strain curve in the elasticregion is defined as the material's stiffness or modulus (in tension, Young's modulus). If thesample is stressed beyond point B, yielding or plastic deformation ensues such that if at point Cthe load were again removed, the sample would return its elastic component of deformationalong line C-C' but not the plastic portion. At C' the unloaded specimen is permanently elongateda distance A-C' and by convention the specimen is said to have a permanent strain, C'; expressedin percent. Loading beyond point C continues to plastically deform the sample until the ultimatetensile strength of bone is reached and failure occurs. Stiffness or modulus is an importantmaterial parameter relating degree of deformation to applied load. As an example, consider thestress-strain curves of three very dissimilar materialsÑ soft metal, glass, and boneÑin Figure 12-5. Metal has the highest stiffness (elastic modulus) and at stresses beyond its yield point exhibitstypical ductile behavior (large plastic deformations before failure) in the nonelastic region. Glassalso has a much higher modulus than bone; however, as a material it readily undergoes brittlefailure, having no discernible nonelastic (plastic) region. Bone has a much lower modulus thaneither metal or glass, but in terms of failure mechanics bone behaves similar to glass, fracturing

in a brittle (low-deformation) mode. Brittle materials fracture into two or more pieces that havevery little permanent plastic deformation and therefore have the potential to be nicely piecedtogether into the original prefracture conformation. In the case of bone, this behavior facilitatesaccurate anatomical fracture reduction and reconstruction using internal fixation. A fracturedmetal plate, in contrast, does not conform to its original prefractured shape, owing to thepermanent plastic deformation that occurred prior to fracture. However, the ductile behavior of the plate is desirable for internal fixation purposes, allowing the surgeon, within limits, toplastically contour the plate to bone without incurring brittle fracture.




FIG. 12-3 (A) Idealized force/elongation plot of the relative mechanicalbehaviors of the tibia and fibula depicted in Figure 12-2. The fibulaexperiences failure at a lower force, Ff, than the tibia, Ft. This graphrepresents the structural parameters of the tibia and fibula. (B) Incontrast, material parameters are represented on a plot of stress versus

strain where the failure loads in A are divided by the correspondingbone cross-sectional areas, thus yielding the relative tensile strengths(stress). Additionally, elongation (delta L) is converted to strain (deltaL/Li). Under such conditions the bone of the fibula and the bone of thetibia (as a material) can be shown to have approximately equivalentstrength (Sf = St).

FIG. 12-4 Stress strain curve for a ductile metal of known cross section. sigma u, is the ultimate strength of the material while sigma y is the yield point that marksthe transition between elastic and plastic behavior. Thearea under the stress strain curve reflects the energy

stored by the material prior to failure. The slope of theplot in the linear elastic region is known as the modulusand reflects the material's stiffness.

FIG. 12-5 Idealized stress strain behavior for three materials:metal, glass, and bone. Note the relative differences in stiffness(slope) and ultimate strength. Bone, although exhibiting someplastic deformation (deviation from linearity), behaves morelike the brittle glass than the ductile metal. (Frankel VH,Nordin M: Basic Biomechanics of the Skeletal System, p. 19Philadelphia, Lea & Febiger, 1980)

The relative stiffness or modulus of bone versus metallic plate is an important consideration infracture healing under conditions of internal fixation. There exists an obvious modulusdifference, or modulus mismatch, between stainless steel and bone. This modulus mismatch willbe discussed more fully in another section; however, the higher stiffness and strength of themetal plate relative to bone in a bone/plate reconstruction result over time in a well-recognizedbony resorption under the plate, so- called stress protection, or plate-induced osteopenia. That is,the bone under the plate atrophies from disuse: the metal plate acts as a stiff bridge protecting theunderlying bone from experiencing forces either axial, bending, or rotational. This phenomenonwarrants precaution when removing metallic plate and screws based solely on radiographicevidence of fracture healing. The "stress- protected" bone, once free of the metal plate, must begradually introduced to weight-bearing loads. Impact forces are to be avoided. Controlled loadsover time allow the bone to remodel and to gradually reassume its initial strength.

FIG. 12-6 Patterns of deformation.




















FRACTURE ETIOLOGYForce is defined as a vector quantity having magnitude and direction. The forces and moments(rotational forces) applied to bone result in loading modes that require understanding to

appreciate the biomechanics of fracture etiology. Bone as a structure may be loaded in tension,compression, bending, shear, torsion, or a combination of these modes. If the magnitude of applied load does not exceed the bone's elastic limits, fracture does not occur, and the bone,elastically deformed, returns to its prestrained state. Shortly beyond the elastic limit, however,catastrophic failure takes place.

The failure mode of bone under circumstances of catastrophic overload is directly related to theloading mode of the bone. That is, from an evaluation of the fracture characteristics, it is possibleto speculate what loading modes produced the fracture. Following is a discussion of loadingmodes, pure and combined, with corresponding clinical examples of failure modes produced byoverload.

TENSIONTensile loading, schematically depicted in Figure 12-6, A, produces an elongation and narrowingof a structure. Maximal tensile forces in the structure are generated on a plane perpendicular tothe applied load, and by definition these are termed normal forces. Consequently, failuretypically occurs along this plane. In bone the tensile failure mechanism has been shown to bemainly one of debonding at bone cement lines and pulling out of osteons. (35) There arerelatively few bones in the body that experience pure tensile forces over their cross-sectionalarea. Most notable are the traction apophyses, including the olecranon process, tuber calcis, andtibial tuberosity, as well as ligamentous attachment sites. Figure 12-7 is an illustration of acommon fracture of the tibial tuberosity produced under conditions of tensile loading.

Because bone as a structural component of the musculoskeletal system must withstand largeaxial loads (both compressive and tensile) to sustain weight bearing and locomotion, it, byadaptation, exhibits greater strength when subjected to tension directed longitudinally versustension directed transversely. This observation is essentially a restatement of Wolff s law andhelps to explain bone's anisotropic mechanical behavior (i.e., varying strength as a function of load direction). Figure 12-8 demonstrates the relative strength of cortical bone tested

FIG. 12-7 Radiograph of an injury produced by tensile forcesimposed on the tibial tuberosity by the quadriceps musculature.

(Courtesy of RB Hohn, DVM) in tension as a function of stressorientation. Clearly, cortical bone loaded longitudinally is stronger,stiffer, and absorbs far more energy to failure than any other loadorientation.

COMPRESSIONCompressive forces on a structure tend to shorten and widen it as shown in Figure 12-6B. As inpure tension, maximal stresses occur on a plane perpendicular to the applied load; however, the








stress distribution and resultant fracture mechanics in compressive failure are often verycomplicated, particularly for an anisotropic material such as bone. Unlike failure in tension,compressive failure in bone does not always proceed along the theoretic perpendicular plane of maximum stress, but rather once a crack is initiated it may propagate obliquely across theosteons(35) following the line of maximum shear stress. Clinical examples of this type of

fracture mode are the compressive fractures commonly seen in vertebral bodies. On occasion,axial loading of long bones may produce compression fractures, particularly at the growth platein young dogs.

SHEARTensile and compressive forces act perpendicular or normal to a structure's surface. In contrast,shear forces act parallel to the surface as shown in Figure 12-6C and tend to deform a structure inan angular manner; thus, squares or rectangles under shear loading become parallelograms. Incontrast to normal forces (tension and compression) where deformation or strain is related tostress by the relationship where:

sigma= E x epsilonwhere sigma=stress, epsilon=strain, E=modulusShear strain is related to stress in terms of angle changetau= G tan alphawhere tau=shear stress,alpha=shear strain angle, G=modulus of rigidity (shear modulus)

Shear loading is described graphically by a load- angle-change plot instead of by a load-deformation plot as in the case of tension and compression. All other considerations, such asstiffness, yielding, and elastic and nonelastic behavior, are interpreted similarly from the plotsregardless of loading mode.

Pure shear fractures are frequently encountered in veterinary orthopaedics; the most common isthe lateral condylar fracture of the distal humerus. Less frequently the tibial plateau, femoralcondyles, glenoid cavity, vertebral bodies, and carpal and tarsal bones are prone to shearfractures. Figure 12-9 is a schematic diagram showing the biomechanics of a lateral condylarfracture of the distal humerus. This fracture typically results from an animal falling or jumpingfrom a height. Owing to the anatomy of the elbow and particularly the distal humerus, axialcompressive forces transmitted up the radius are imparted largely (80%) to the lateral condyle,producing shear forces in the intercondylar and epicondylar areas. In Figure 23-2A a radiographshows a typical lateral condylar fracture resulting from shear forces exceeding the shear strengthof bone in the intercondylar and epicondylar areas.

FIG. 12-8 Anisotropic behavior of cortical bone specimensmachined from a human femoral shaft and tested intension. The orientation of load applicationÑlongitudinal(L), tilted 30¡ with respect to the bone axis, tilted 60¡, andtransverse (T)strongly influences both the stiffness and theultimate strength. (Frankel VH Nordin M BasicBiomechanics of the Skeletal System, p. 22 Philadelphia,Lea & Febiger, 1980)








FIG. 12-9 Biomechanics of a lateral humeral condyle fracture. Forcestransmitted proximally along the radius impact largely on the lateralhumeral condyle, thus creating shear forces in the intercondylar andepicondylar areas, where failure is prone to occur.

It is important in describing fracture etiology to understand the mechanical behavior of boneunder these three primary modes of loading. It has been shown that cortical bone strength isstrongly dependent on the mode of loading, being strongest in compression and weakest in shear(Fig. 12-10).(77) Thus, bone strength is a function not only of load orientation within any givenmode as shown in Figure 12-8, but it is also a function of loading modeÑtension, compression,or shearÑat a given load orientation. This observation helps to explain why fracture lines do notfollow precisely the line of maximum stress (e.g., in compressive failure in which maximumstress occurs perpendicular to applied load). Rather, owing to mechanical anisotropy cracks maypropagate obliquely along lines of maximum generated shear stress and reduced bone strength.This consideration applies also to failure under conditions of bending and torsion yet to bediscussed.

BENDINGBending is a loading mode schematically illustrated in Figure 12-6D and results in the generationof maximum tensile forces on the convex surface of the bent member and maximum compressiveforces on the concave side. Between the two surfaces, that is, through the cross section of themember, there is a continuous gradient of stress distribution from tension to compression (Fig.12- 11). An imaginary longitudinal plane corresponding to the transition from tension tocompression, approximately in the center and normal to applied force, is designated the neutralsurface. Along this surface there is theoretically no tensile or compressive load on the material.Another useful designation is the neutral axis, which is the line formed by the intersection of theneutral surface with a cross section of the beam, perpendicular to its longitudinal axis (Fig. 12-11).

FIG. 12-10 Ultimate stress for human adult cortical bonespecimens tested in compression, tension, and shear. Forcomparative purposes the shaded area represents the ultimatestress in tension and compression for human adult cancellousbone with a density of 0.35. (Frankel VH Nordin M BasicBiomechanics of the Skeletal System. Philadelphia, Lea &Febiger, 1980; data from Reilly D, Burstein A: The elastic andultimate properties of compact bone tissue J Biomech 8:393,1975)

FIG. 12-11 (Top) A beam subjected to pure bendingshows the relative orientation of the neutral surfaceand the neutral axis. (Bottom) Distribution of stresses around the neutral axis in a transversesection of the beam subjected to bending. Tensilestresses are maximum on the top surface, and































compressive stresses are maximum on the bottomsurface of the beam The neutral surface theoreticallyexperiences no tension or compression.

During normal function bone is subjected to large bending forces both intrinsic and extrinsic

(termed moments). The act of locomotion, for example, results in alternating tension andcompression on the cortex of weight-supporting bones during the gait cycle.(19,35) Theintroduction of large extrinsic forces (e.g., automobile trauma) perpendicular to the diaphysis of long bones may generate enough tension on the convex surface of the bent bone to exceed itsinherent tensile strength, resulting in crack initiation and failure. Clinically, fractures producedby bending forces are commonly transverse or short oblique, as shown in Figure 12-12. Themechanism of failure in bending is one of crack initiation at the point of maximum tensile stresson the convex (tension) surface of bone with crack propagation along a line of maximum tensilestress or minimal material strength (e.g., in shear) resulting in transverse or short obliquefractures, respectively. Because mature healthy bone is stronger in compression than in tension,failure usually begins on the tension surface. In very young animals or severely osteoporotic

bone, however, folding or buckle fractures are sometimes noted on the concave or compressionside of the bone, indicating failure in a compressive mode subsequent to bending.

FIG. 12-12 Radiograph of a short oblique long-bone fractureprobably produced by bending forces imposed on the midshafthumerus.

Structurally, in a bending mode of loading, bone strength and stiffness are dependent not only on

cross-sectional area as in tension and compression but also on the arrangement or distribution of bone mass about the neutral axis (shape). This strength parameter is termed the area moment of inertia and is an important concept in understanding the strength of a specific shape or geometryunder conditions of bending. For example, it is intuitive that a 2" x 4" piece of lumber is"stronger" in bending when placed on its edge (2" side) than on its flat (4") side, yet cross-sectional area remains constant. From the beam theory, formulas have been derived to expressarea moment of inertia as a function of geometry. The formula to compute the area moment of inertia, I, for a rectangular cross section is

I= base(height)^3 /12

From Figure 12-13 one can appreciate, therefore, that a 2" x 4" on its edge has an area momentof inertia four times greater than on its side and accordingly demonstrates a fourfold increase inrigidity. More simply put, the area moment of inertia takes into account the fact that in bending,a structure gets stronger (and stiffer) as its mass is moved further from its neutral axis. Inengineering applications this concept is demonstrated nicely by the design of the "I" beam, whichaffords maximum resistance to bending with minimum weight. Long bone, in its tubular shape,is aptly designed to uniformly resist bending in all directions and in addition has its mass located








circumferentially at a distance from the neutral axis, thus providing a high area moment of inertiaand high resistance to bending.

FIG. 12-13 The area moment of inertia, 1, for a rectangularbeam loaded in bending is I = bh^3/ 12. The calculation for

a 2" x 4" beam yields a moment of inertia approximatelyfour times greater when the beam is loaded on its edge thanon its side. This corresponds to a fourfold increase inrigidity. Moreover, the maximum bending stress in thecross section of the beam considered is proportional to thedistance from the neutral axis and inversely proportional to1, indicating that the 2" x 4" beam on its edge whencompared to loading on its side can withstand roughlytwice the load necessary to cause failure.

TORSION

Torsional loading as depicted in Figure 12-6C is a geometric variation of shear and acts to twist astructure about an axis (the neutral axis). The amount of deformation is measured in terms of shear angle, alpha. As in bending, in which maximum tensile and compressive stresses occur onthe surface and distant from the neutral axis, torsional loading produces maximum shear stressesover the entire surface, and these stresses are proportional to the distance from the neutral axis(Fig. 12-14).

The fracture mechanics in torsional failure are more complicated than those in any of the otherloading modes previously described. A material under torsional loading experiences maximumshear stresses on planes perpendicular and parallel to the neutral axis, while maximum tensileand compressive stresses are generated normal to each other and on a diagonal to the neutral

axis. This is more clearly demonstrated in Figure 12-15 where the square finite element drawn onthe surface of the cylinder undergoes shear-type deformation with torsional forces applied to thecylinder. The diamond-shaped element experiences a deformation in torsion analogous to simpletension and compression; that is, it elongates and narrows, with maximum tensile stresses actingon a plane perpendicular to the axis of elongation and maximum compressive stresses orthogonalto this. Considering the stress distribution in areas other than the principal axes of tension andcompression, it is apparent from Figure 12-16 that on planes perpendicular and parallel to theneutral axis maximum shear stresses are manifested in this material. In torsion, then, as in otherloading modes, the location of crack initiation and the direction of its propagation are dependenton the inherent strength of the material in any given loading mode and on the magnitude of theimposed stresses within the material. In dog bone subjected to pure torsional loading, it has been

suggested that failure begins with crack initiation in a shear mode,(35) that is, parallel to theneutral axis, followed by crack propagation generally along the line of maximum tensile stress(30¡ to the neutral axis). The net effect of this fracture mechanism is to produce a so-called spiralfracture of the long bone as shown schematically in Figure 12-17. Figure 12-18 is a radiographshowing a spiral fracture of a humerus as observed in clinical practice.











FIG. 12-14 Cross section of the cylinder loaded inFigure 12-6, E in torsion shows the shear stressdistribution about the neutral axis Shear stressincreases as a function of distance from the neutralaxis.

FIG. 12-15 Torsional deformation of a square-shaped and adiamond-shaped element drawn on the surface of a cylinder.The elongation of the diamond-shaped element with torsionindicates that maximum tensile stresses are acting on a planeperpendicular to the axis of elongation and, conversely, thatcompressive stresses act orthogonal to this. The squareelement undergoes shear-type deformation similar to thatshown in Figure 12-6,C.

FIG. 12-16 Maximum shear stresses on thesurface of a cylinder subjected to torsional

forces occur on planes perpendicular andparallel to the neutral axis. Spiral fracturesof bone are produced by torsional forcesapplied to the overall structure; failure of the material, however, typically occurs intension along the line of maximumgenerated tensile stress.

FIG. 12-17 Schematic illustration of a two-piece spiral fracture of thefemur drawn from a radiograph of a clinical case Note the orientation of the fracture line relative to the long axis of the bone. In this example bothshear and tensile failure modes are represented.

As in bending, in which strength and stiffness of bone are determined by the area moment of inertia, that is, the size and shape of the bone, the analogous structural quantity in torsionalloading that takes into account size (cross-sectional area) and shape (distribution of bone aboutthe neutral axis) is the polar moment of inertia. This quantity provides an explanation for theobserved "structural" strength of healing fractures in which abundant callus has formed a cuff around the fracture ends. Obviously fracture callus does not have the material strength of organized lamellar bone; however, the net effect of large cross-section and callus distributiondistant from the neutral axis gives fracture callus a polar moment of inertia approaching the

strength and stiffness of whole intact bone. The polar moment of inertia also explains whycortical bone at the isthmus of long bones (small diameter) must perforce be thicker than in thewider metaphyseal areas to produce equivalent resistance to torsional and flexural loading.

FAILURE UNDER COMBINED LOADINGTension, compression, shear, bending, and torsion represent simple and pure modes of loading.Examples of failure under these loading modes have been presented. In clinical practice,




















however, fractures encountered are more commonly a product of a combination of theaforementioned modes. This is not surprising when one considers that the mode of loading isdetermined by the direction of load application and that in the case of bone fracture produced bytrauma (e.g., automobile) there is virtually no constraint on applied load orientation (ormagnitude).

FIG. 12-18 Radiograph of a spiral fracture of the humerus shows thespiral fracture surface and a fissure extending proximally along theline of maximum tensile stress.

ENERGY TO FAILUREAs shown in Figure 12-4, the area under the stress-strain or load deformation curve correspondsto the energy absorbed by the bone while undergoing deformation. Since bone behaves largelylike a brittle material, exhibiting very little permanent plastic deformation to failure, most of thisabsorbed energy is returnable upon unloading. When bone is loaded to failure, however, thestored energy is released or dissipated at a very rapid rate through the formation and propagationof one or more cracks. The number and pattern of cracks formed depend largely on the rate atwhich load is applied. Bone has been shown to have a higher modulus (stiffness) and to absorbmore energy to failure the more rapidly it is loaded,(80) that is, it is stiffer and tougher. A singlecrack, however, has a finite threshold energy for initiation and a finite capacity to dissipatestored or applied energy. Thus, under conditions of high loading rate, if the stored energy in thestructure exceeds that which can be dissipated via the formation of one crack, multiple crackswill form and energetically less favorable fracture mechanisms may initiate. This situationresults clinically in fracture comminution. Stated in another way, bone has a finite capacity toabsorb energy that increases significantly with load rate. When the energy extrinsically impartedto bone (kinetic energy) exceeds the energy- storage capacity of bone, fracture occurs. Kinetic

energy is defined by the formula

KE= 1/2mv2 where m = mass v = velocity

From this formula it is clear that the effect of increasing load rate, that is, the velocity, plays abigger role in determining the ultimate fracture (and fracture potential) than does mass alone.(This point will be discussed in the chapter on ballistics.)

Fractures are arbitrarily grouped into three general categories based on the energy required toproduce them: low-energy, high-energy, and very high energy fractures. An example of a typicallow-energy fracture would be the lateral humeral condyle fracture that results when a Yorkshire

terrier falls from its owner's arms. High-energy fractures are commonly observed followingautomobile trauma (Fig. 12-19), and very high energy fractures are associated exclusively withgunshot injuries produced by missiles having high muzzle velocity








FIG. 12-19 Radiograph of a high-energy fracture shows markedcomminution resulting from trauma sustained in an automobileaccident.

FIG. 12-20 Radiograph of a very high energy fracture produced by abullet fired from a gun at high muzzle velocity. The energy of theprojectile imparted to the soft tissue and bone upon impact results inextensive tissue destruction and very fine bony comminution. associatedexclusively with gunshot injuries produced by missiles having highmuzzle velocity (Fig. 12-20)

FATIGUE FRACTURESFatigue fractures are infrequently encountered in the practice of small animal orthopaedics. Theyare, however, a common occurrence in human and equine practice and also in certain dog

sporting events, such as dogsled racing, for which dogs must be trained to the limits of theirendurance.

Fatigue fracture is a phenomenon observed in many materials systems including bone. Incontrast to the bone fractures previously mentioned, in which failure followed static loading of bone beyond its ultimate stress, fatigue fractures result from repetitive loading of bone atmagnitudes below the ultimate strength of bone. Clinically, fatigue fractures occur afterprolonged periods of strenuous activity in which cyclic loads coupled with muscular fatigue(exhaustion) are predisposing factors.

FIG. 12-21 S-N curves for an idealized metal and for

cortical bone show a marked difference in fatigue behavior.At stresses below the endurance limit, the metal can becycled endlessly without experiencing failure. Deadcortical bone, however, has been shown to be susceptible tomicrodamage and ultimate fatigue failure at small loadmagnitudes well below the ultimate strength of bone sigmau and therefore is thought not to demonstrate arecognizable endurance limit. (Carter DR The FatigueBehavior of Compact Bone. PhD dissertation, StanfordUniversity, 1976)

The fatigue behavior of a material is classically represented on a plot of the peak stress per cycle(S) versus number of cycles (N) to failure as shown in Figure 12-21. The quantity sigma u,represents the ultimate strength of the idealized material under a load frequency of N = 1. Pointson the curve represent the S-N conditions that produce fracture of the material. At low frequencyand high stress, fracture occurs by mechanisms other than fatigue, that is, failure due to repetitiveloading of a material beyond or near its yield point. At high frequency, however, and well belowa materials yield point, fracture occurs by fatigue mechanisms, namely, microcrack formationand crack coalescence. The dotted line corresponds to a so-called endurance limit that many












materials display. A material can be cycled virtually endlessly at stresses below this limit. Carterand Hayes,(20) however, have studied fatigue properties of cortical bovine bone and havedemonstrated by flexural testing that microdamage is prevalent even at low-frequency cyclingand at loads well below the ultimate flexural strength of bone. Based on these results, Carter(18) suggested a mathematical relationship to predict the number of cycles to failure under fatigue

conditions. Interestingly, the formula does not make provision for an endurance limit as occurs invarious other inanimate materials. This information would imply that dead bone is susceptible tomicrodamage and fatigue failure even under small-load magnitudes given adequate cycling. Inliving bone, however, it is theorized that microdamage may be a stimulus to bone remodelingand that catastrophic fatigue failure occurs only when the rate of damage outpaces the rate of biologic repair (7,20,25,28,37)

BIOMECHANICS OF FRACTURE REDUCTIONBiomechanical principles of fracture reduction are simple but, unfortunately to date, are not

quantifiable. Each surgeon with time develops an individualized assortment of fracture reductiontechniques specific to fracture type. Whether the reduction is performed open or closed,manually or with skeletal traction, with or without instruments, the objectives remain the same:

1. The fractured bone ends and fragments must be brought into close enough proximity tooptimize the fracture-healing process.2. The reconstructed fracture must approximate normal anatomy well enough to provide foroptimum function after healing.3. The preceding must be accomplished with minimal additional trauma to vital structures andsurrounding tissue.These requirements necessitate a mechanical means of applying force either remotely or locallyto mobilize the fracture ends and move them into acceptable orientation. The resultant force toachieve reduction is largely tension (traction) along the axis of the long bone and must besufficient to overcome gravitational forces (limb weight); forces of muscle contracture;hydrostatic forces due to edema; and in the case of long-standing fractures, forces due togranulation tissue and fibrous callus at the fracture site. The mechanical effect of edema as animpediment to fracture reduction is often not fully appreciated. Postfracture edema andhematoma in the course of achieving a hydrostatic equilibrium fill interstitial spaces and createfluid-filled voids along tissue planes surrounding the fracture site. The effect is to impart lateralforces circumferentially to the soft tissue overlying the fracture. The process stops, that is,equilibrium is achieved, when the hydrostatic pressure from edema is counteracted by tension inthe walls of soft tissue compartments and skin. The expansile forces act to shorten the fracturedextremity and resist reduction, thereby freezing the fracture orientation in its shortened position(Fig. 12-22). An attempt at fracture reduction during the phase of edema is both difficult andhazardous. Without an option for protracted skeletal traction, it is often judicious, if notnecessary, to apply a temporary compression wrap to the limb for 24 to 48 hours to combat thelateral expansile forces and minimize edema before fracture reduction is attempted.


















Forces necessary to overcome muscle contracture and fibrous callus increase with time(days/weeks), and therefore to facilitate fracture reduction it is advisable to attempt a reductionas early in the postfracture period as is clinically feasible. (See Chapter 14.)

FIG. 12-22 Idealized model of fracture edema leading to

compression and shortening of the extremity. To achieve fracturereduction in the presence of significant edema, forces must beapplied to overcome the lateral expansile forces of edema fillingtissue compartments and interstitial spaces. Alternatively acompressive wrap can be used prior to fracture reduction tominimize or reverse the extent of edema formation.

For closed nonsurgical realignment of long-bone fractures, reduction entails grasping through theskin, the joints, and bony prominences proximal and distal to the fracture. The tension necessaryto effect reduction at the fracture site can be achieved only by exerting compressive, tensile, andshear forces on the soft tissues overlying or remote from the fracture site. Soft tissues, however,

like bone, are materials having yield points and failure limits, and the surgeon must therefore usecare not to exceed these ultimate biologic limits to avoid additional tissue damage in reducing afracture. Typically biologic strength is a small percentage (10%) of the tissue's material strength,and therefore considerable iatrogenic tissue damage in the process of fracture reduction is notinconceivable.

The most effective closed technique for achieving reduction of transverse or short oblique long-bone fracture is "toggling," as shown in Figure 12-23. It must be recognized, however, thatdigital pressure exerted over sharp fracture ends creates very high concentrations of stress thatare potentially injurious to soft tissues and may result in the creation of an open fracture.Accordingly, digital pressure over sharp fracture ends should be avoided if possible. If attempts

at closed reduction yield poor results, open surgical reduction is indicated. This topic is discussedmore fully elsewhere (Chapter 16). The mechanical act of open reduction necessitates anassortment of instruments, the purpose of which is to allow the surgeon to grasp and localizedistraction forces in proximity of the fracture ends. Ideally a bone-holding forceps should satisfythe following requirements: (1) securely grasp bone via self-retaining mechanism withoutcausing further comminution; (2) cause minimal soft tissue injury and periosteal stripping duringapplication; (3) facilitate anatomical realignment of fracture fragments; and (4) have the capacityto maintain this reduction without obstructing the application of primary internal fixation.Obviously no single bone-holding instrument can satisfy these requirements under all fractureconditions, and therefore an assortment is indicated. At times it may be necessary to useadditional aids in maintaining reduction, such as full cerclage wires (which, if made of stainless

steel 316L, may be left permanently in place) or temporary full cerclage nylon bands (normallyused as collars for bundles of electrical wires; the latter must be removed prior to closing).












FIG.12-23 Illustration of the method of "toggling" to achievefracture reduction. Bone fragments must be placed in correctrotational orientation and angled to provide edge-to-edge corticalcontact. Gradual straightening of the bone is accomplished by digitalpressure applied away from sharp fracture ends.

The surgeon should be aware of some mechanical limitations to achieving an adequateintraoperative reduction. First, bone, having limited ultimate strength, will undergo furtherfracture if excessive force, either in grasping the bone or in physically distracting fragments, isimprudently exerted. Second, the use of orthopaedic instruments gives the surgeon a largemechanical advantage that must be used judiciously in accordance with the ultimate biologicstrength of surrounding vital structures (e.g., nerves and vessels). Exceeding the biologic limitsof these tissues can result in purposeless surgery. Finally, forces necessary to achieve reductionmay exceed the surgeon's inherent physical strength. This situation can potentially arise fromfractures in very large dogs, fracture of long-standing duration (2 weeks), or fractures havingrelatively inaccessible location, such as ilial shaft fracture.

BIOMECHANICS OF FRACTURE FIXATIONIn veterinary orthopaedics there exist two major objectives of fracture fixation: first, fracturefixation must provide for early if not immediate ambulation and weight bearing, and second, itshould optimize the bone-healing process such that the need for fracture fixation is supplanted ina minimum period of time by recovery of strength and stiffness of the healed bone.Unfortunately, these two objectives at times run counter to each other. For example, the rigidityof fracture fixation achieved through the use of internal fixation with plate and screws satisfiesthe first objective and provides for early weight bearing. However, the overwhelming strengthand stiffness of the device is thought to shield the bone from normal weight-bearing stresses andtherefore short circuit the biomechanical stimuli necessary to attain optimum strength andrigidity of bone. It is thus likely that the choice of fracture-fixation methodÑinternal, external, ortransfixationÑmust represent a trade-off of these two conflicting objectives. The choice,however, is not very clear-cut. The surgeon in the morass of confusing animal experimentation,subjective reporting, and as yet poorly understood fracture-fixation biomechanics, by necessitypractices orthopaedics with those fracture-fixation techniques that work best in his hands.Fortunately, encouraging developments in biomechanical experimentation are providingquantitation and means of optimization of the fracture-healing process.

As discussed, the specific mechanical aims of fracture fixation regardless of method are to

provide stability of fracture reduction, namely, maintain axial alignment and prevent rotation,and to support the fracture in this manner until the completion of fracture healing.Correspondingly, the topic of biomechanics of fracture fixation can be divided into two sections:biomechanics of fracture stability in the immediate postfixation phase and biomechanics of healing relative to method of fracture fixation. The discussion to follow will attempt to coverpertinent information in each of these areas.




EXTERNAL IMMOBILIZATIONThe objective of external immobilization is to stabilize fracture fragments internally byapplication of rigid material externally. The process necessitates interposing soft tissue betweentwo relatively rigid materials, bone and splint. The fracture stability achieved with externalfixation, unlike that of internal fixation, is not rigid per se but relative, that is, dictated by the

limits of compressibility or strain tolerance of interposed soft tissue. Thus, by definition externalimmobilization in the immediate postfixation period can at best approach but not equal thestability of rigid internal fixation or transfixation. Nevertheless, fracture healing can proceedtoward so-called secondary bony union if the applied external fixation provides limits to fracturefragment displacement compatible with specific biologic processes of repair, particularlyrevascularization.

Owing largely to technical difficulties in biomechanically evaluating fracture stability underconditions of external immobilization, this topic has received surprisingly little attention in theliterature. Moreover, the necessities of veterinary orthopaedics (in contrast to humanorthopaedics) dictate that all forms of external fixation be functional, that is, weight bearing.

Therefore the information extractable from human literature is reduced even further. Two areas,however, deserve mention: functional bracing and the material properties of casts.

Functional bracing was first introduced in the late 1960s by Sarmiento,(81-83) a humanorthopaedist. The original concept applied to tibial fractures incorporated a below-the-knee castthat hypothetically transmitted ground reaction forces to the straight patellar tendon, bypassingthe fracture. Further investigation, however, including accounts from patients who experiencedpressure over the calf area during ambulation, pointed toward a more plausible mechanism.Plastic braces were instrumented with pressure transducers and evaluated during ambulation.Results revealed significantly greater pressures over the gastrocnemius muscle group than thetibial tuberosity, suggesting that perhaps the original hypothesis was invalid. Additionally, theplastic brace itself was found to carry only 15% of the axial load during weight bearing,indicating that the constrained bone and soft tissue were withstanding greater than 80% of theground reaction forces. Further studies of anesthetized patients having tibial fracturesdemonstrated a 75% reduction in bone overriding of braced fractures versus nonbraced fractureswhen subjected to 25-pound axial loads. These findings led Sarmiento(49,84,85) to postulate amultifactorial mechanism operating to prevent shortening and angulation during functionalbracing. First, the calf musculature, owing to its incompressible nature and conical shape, whenconstrained by a conforming plaster brace develops large hydrostatic pressures during weightbearing that act to prevent shortening. Second, the elastic properties of soft tissue structures inthe area of the fracture prevent excessive overriding. Third, the interosseous membrane betweenthe tibia and fibula further resists fracture displacement.

The principle of functional bracing has been adapted for use in tibial fractures in dogs byNunamaker. A two-part thermoplastic splint consisting of an cranial half cast from stifle to toesand a cranial conforming brace from stifle to just above the hock is applied using elasticbandage. Experimental and clinical results to date for fractures involving the proximal two-thirdsof the tibia are encouraging. Specific techniques of splint application will be covered in Chapter15.










For reasons of economy and acceptable performance, plaster casts find widespread application inveterinary orthopaedics. Biomechanical studies on the properties and stresses in orthopaedicwalking casts have been carried out by Schenck and co-workers.(88) Although their workinvolves lower-leg walking casts for human application, the results are in part applicable toweight-bearing casts in animals. Several findings are of interest. First, the ultimate strength of

plaster is finite and in tension is one third that in compression (700 psi versus 1800 psi). Thiswould suggest in general that areas of a cast subjected to large tensile stresses or large resultantmoments about joints such as occur during walking may require buttressing with extra layers of plaster. Examples in the dog would include the caudal surface of the carpus in an extension castof the forelimb, or in the case of a lower hind-leg cast, the cranial surface of the hock, which inhumans at least is prone to failure owing to high tensile forces in plantar flexion.

Second, the strength of plaster casting material is dependent on setting and drying time. Schenckand co-workers(88) have shown that after 24 hours (the recommended drying time), a 1/4"plaster cast will achieve only about one third of its ultimate strength.. The authors suggested atleast 48 to 60 hours drying time before the casts are subjected to weight-bearing forces.

To the veterinary orthopaedist, these findings would indicate that plaster casts be constructedwith optimum strength using the minimum amount of materials to facilitate drying, lightness of weight, and early weight-bearing. A heavy, cumbersome cast is just as contraindicated as onethat is too weak. To facilitate the drying process the fresh plaster cast should not be wrappedwith occlusive dressing (including adhesive tape and elastic bandage) for a period of at least 24hours (for 1/8'' cast). Until the cast is completely dry, it is recommended that the patient be cagerested.

EXTERNAL SKELETAL FIXATION

External skeletal fixation (also termed transfixation or fixateur externe) is a fixation methodconsisting of multiple percutaneous, transcortical pins proximal and distal to the fracture siteincorporated into a surrounding external frame. The system facilitates easy management of fracture-associated, soft tissue injuries while providing adequate skeletal fixation without thephysical presence of metal at the fracture site. Accordingly, external skeletal fixation is suitablefor the treatment of compound fractures and infected nonunions as well as for stabilization of joint fusions, osteotomies, and limb-lengthening procedures. Owing to renewed interest in thismethod by human orthopaedists, several external fixators have been commercially introduced invarious frame configurations (Fig. 12-24). High cost, however, has largely precluded theirpractical use in veterinary orthopaedics. The system most frequently used in veterinary practicehas been the Kirschner-Ehmer (K-E) apparatus. The K-E instrumentation is readily adaptable forapplication at various skeletal locations (tibia, femur, humerus, radius, and ulna) and can, withpractice, be constructed into appropriate frame configurations, the more common of which areshown in Figure 12-25. The type and location of fracture and the condition of soft tissues will of course determine the final frame configuration.














FIG. 12-24 Common geometric shapes used in the designand application of external fixators: (a) The unilateralconfiguration; (b) the bilateral configuration; (c) thequadrilateral configuration; (d) the triangular configuration(half-pin); (e) the triangular configuration (full- and half-

pin); (f) the semicircular configuration; (g) the circularconfiguration. (Chao EY, Kasman RH, An KN: Rigidityand stress analyses of external fracture fixation devices: Atheoretical approach. J Biomech 15:971, 1982)

As with other fixation methods, the successful application of external skeletal fixation dependson an understanding of the mechanics of fracture reduction and fixation, as well as an awarenessof the structural integrity of the bone-fixator composite. Research both theoretic andexperimental is ongoing and has provided useful information in this regard.(22) The clinicalperformance of external skeletal fixation is dependent on fixator stiffness, which has been shownto be related to the geometric arrangement of pins, including their direction of orientation and

their diameter, the length of pins between bone and clamp, the distance between clamps, thelengths of sidebars, and the design and system of assembly of the frame. Many of these factorsare determined by the surgeon at the time of fracture fixation, and therefore surgical techniquehas a profound influence on both the stability of the fixator and motion at the fracture site.(45)

FIG. 12-25 Configurations of Kirschner-Ehmerapparatus commonly used in veterinary practice. (A)Double-clamp (half-pin) configuration. (B) Singleconnecting-bar (half-pin) configuration. (C) Doubleconnecting-bar (full-pin) configuration. (D)Triangular (half-pin) configuration.

As discussed in Chapter 16, the two most common fixator configurations used in veterinarypractice are the half frame (unilateral configuration; Fig. 12-25, A) and the full frame (orbilateral configuration; Fig. 12-25, B). The K-E half frame is most similar to the Hoffmann half-frame system, the mechanics of which have been extensively studied.(13) Lindahl,(54,55) in aseries of investigations in the early 1960s, examined the rigidity of various kinds of osteosynthesis for transverse and oblique fractures; he concluded that the rigidity of the half-frame configuration was unsatisfactory from a clinical view- point. In contrast, Burny and co-workers(13) conducted theoretic and experimental studies to determine the optimum conditionsof half-frame fixation and concluded that the connecting sidebar length should be minimized;clamps must be as close to the bone as possible; and better stability is achieved with three pins

than two, maximally spaced in each fragment. Under optimum conditions, however, the half-frame fixator provides "elastic" rather than rigid skeletal fixation. The authors contend that this"controlled mobility" acts to enhance clinical fracture healing through exuberant periosteal callusformation.(13)

Although it is common practice in veterinary orthopaedics to use two or more angled pins ineach fracture fragment in conjunction with the half-frame K-E, the Hoffmann apparatus usesparallel transfixation pins, and therefore a comparative assessment of fixator rigidity is not
































possible. Minimally, the angled pins would seem to act as a desirable impediment to theapparatus being pulled out by the animal. In this vein, Evans(31) has demonstrated improvedrigidity of the Oxford apparatus by angling the outer pins inward by 20¡, while Boltze(8) angledand elastically bent the pins in opposite directions in each proximal and distal segment andobtained greater fixator stiffness. In general, however, the unilateral frame (half-pin) has been

shown to yield only 20% to 50% of the fracture stiffness obtained by a standard full-pin bilateralconfiguration,(14,24,33) and therefore its use has significant mechanical limitations.

The full-frame external fixator (bilateral configuration) uses through-and-through pins andprovides superior bone-fixator stability as well as improved longevity for prolonged clinicalapplication. By manipulating pin size and pin number, a spectrum of fracture stabilities can beachieved from elastic to rigid. Also, since the full frame is symmetric in configuration, it isconducive to applying either compression or distraction forces to the fracture ends.

The relative rigidity of the full-frame external fixator depends on several important geometricvariables. The number of pins in proximal and distal fracture segments has been shown to

influence fixator stiffness directly, with three pins per segment yielding 50% to 100% morestiffness in compression, bending, and torsion than two pins per segment.(22) Interestingly, theaddition of more pins (more than four) per segment results in an insignificant increase in overallfixator stiffness, particularly in the most critical anteroposterior (AP) bending mode. Pinplacement within each pin group is also important. In general, a comparison of various pinnumbers used and their placement in bone has demonstrated that fixator stiffness can beimproved by increasing pin separation within each group; minimizing pin length; and reducingpin-group separation.(9-11,21,23,58) The determination of optimum pin separation, however,must take into account the obvious regional differences in bone stiffness. For example,Egkher(29) has suggested that better overall stability may be achieved by firm placement of pins"closer" together within diaphyseal cortical bone, away from the thinner and weaker cortical andcancellous bone of the metaphysis. Increasing the transfixation pin diameter can greatly improvethe rigidity of the composite bone-fixator. For example, a 2-mm increase in pin diameter (from 4mm to 6 mm) results in a fivefold increase in rigidity of a standard Hoffmann-Vidal frame(bilateral).(24) This is attributable to the area-moment inertial properties of the larger pin; that is,rigidity is a function of the fourth power of the diameter. Clearly, however, pin diameter must bebalanced against the weakening (stress raiser) effect of the associated hole in the bone.Optimally, pin diameter should not exceed 30% of the bone diameter. (11) The diameter andstiffness of the sidebars are typically much greater than those of the transfixation pins. Thesidebars, therefore, are not a limiting factor in the overall fixator rigidity; however, to minimizethe total weight of the external fixation apparatus the surgeon should select the minimumdiameter and length of rod to achieve an optimum combination of strength and weight.

FIG. 12-26 Parametric diagrams illustrate the effects of geometric variables on the stiffness property of thestandard Hoffmann-Vidal external fixation device. (d, pindiameter; h, pin separation; l, sidebar separation [pinlength]; s, pin group separation; n, pin number) (Chao EY,Kasman RH, An KN: Rigidity and stress analyses of external fracture fixation devices A theoretical approach. J




































Biochem 15:971 1982)

FIG.12-27 Overall relative stiffness comparison for variousfixators in the tibia. (Chao EY, Kasman RH, An KN Rigidityand stress analyses of external fracture fixation devices: Atheoretical approach. J Biomech 15:971, 1982)

The content of the foregoing discussion has been neatly summarized by Chao in an analysis of the rigidity and stresses in external fixation devices by means of the finite element method.(23) Figure 12-26 shows parametric plots of the effects of geometric variations on bone fixationstiffness in axial compression, torsion and AP and lateral bending. Although these theoretic plotswere derived using the standard Hoffmann-Vidal (bilateral) configuration as a model, the resultscan be applied to the stiffness characteristics of a full-frame (bilateral) K-E

apparatus with reasonably good approximation. In Figure 12-26, C, it is important to note the

reduced frame stiffness when loaded in the AP mode of bending (< 160 N/cm). This relationshipwas experimentally substantiated in a comprehensive comparative study of various fixatordesigns conducted at the Mayo Clinic.(58) Figure 12-27 summarizes these findings along withsimilar results from the standard Hoffmann,s4 the Hoffmann- Vidal,(22) and the Oxfordfixator.(31) It is clear from Figure 12-27 that all external skeletal fixation devices tested arerelatively weak under AP bending conditions.

Chao, in a comparative study, evaluated the relative fracture stiffness of prepared tibias fixed byeight-hole AO compression plate or common Kuntscher nail and compared these findings withresults obtained from similar loading conditions of the Hoffmann-Vidal fixator in a plastic bonemodel.(24) For comparative purposes the results appear in the compound diagram shown in

Figure 12-28. It is noteworthy that the composite stiffness characteristics of external fixators andinternal fixation devices differ markedly. The plate demonstrates superior stiffness behavior inall but the "bending open" loading mode, whereas the nail, as might be expected, performsrelatively well in bending, AP, and lateral, but has minimal ability to withstand axial loading andtorsion. The very rigid eight full-pin quadrilateral fixator (8 F-pin, Hoffmann-Vidal), except inAP bending as discussed, exhibits rigidity approximating plate fixation. The six full-pin (6 F-pin,Vidal-Adrey) fixator, however, has significantly reduced stiffness, yet in axial and torsionalloading is still superior to nail fixation. Not surprisingly, the six half-pin fixator (6 H-pin,Hoffmann-Vidal) demonstrated the least stiffness of all fixation systems evaluated andaccordingly has been termed "elastic external fixation."(13)

FIG. 12-28 Comparison of fracture fixation stiffnessproduced by internal and external fixation devices: 8 full-pin (8F-PIN) Hoffmann-Vidal splint; 6 full-pin (6F-PIN),Vidal-Adrey splint; 6 half:pin (6H-PIN) Hoffmann-Vidalsplint. (Chao EY8, Pope MH: The mechanical basis of external fixation In Seligson D Pope MG (eds): Concepts inExternal Fixation, p 13. New York Grune & Stratton, 1982)































INTERNAL FIXATIONInternal fixation in veterinary orthopaedics is readily divided into two distinct treatmentmodalities: intramedullary (IM) fixation (pins, wires, and nails) and screw and plate fixation. Theformer category, though little studied in a biomechanical sense, is widely used on an empiricalbasis in veterinary orthopaedics. Screw and plate fixation, on the other hand, has received more

complete investigation, yet because of technical complexities in application and the increasedcost has not received widespread use in veterinary practice.

INTRAMEDULLARY FIXATIONIM fixation as applied to small animals is limited to four basic techniques: single IM Steinmannpin; single IM Kuntscher nail; single IM Steinmann pin with hemicerclage wiring; and multipleIM Steinmann pins with full cerclage wiring. To date few reports appear in the literatureaddressing the biomechanical principles of IM fixation.(3) Moreover, none of the reports pertainto IM fixation using Steinmann pins alone or in combination with antirotational wires. In a studyby Allen and co- workers(3) focusing on IM rod fixation, the most important mechanicalcharacteristics determining device performance of rod fixation in humans were identified as

bending strength, bending rigidity, and torsional rigidity. In this report bending strength andbending rigidity represent "device" properties that, if exceeded either by excessive static load orfatigue, result in implant failure. By contrast, in veterinary orthopaedics, failure of IM fixationrarely occurs as a result of material strength limitations. Rather, failures more commonly stemfrom inadequate torsional rigidity; that is, the fracture site becomes rotationally unstable whilethe rod remains intact. Allen and co-workers have shown that a Kuntscher nail, when optimallyplaced in a reamed cadaver femur and torsionally tested, has only about 15% of the torsionalrigidity of the contralateral unosteotomized side.(3) Even the best designed fluted femoral IMrod yielded a torque capability of only one third the control side. Mensch and co-workers,(59) using similar methodology, added polymethylmethacrylate (PMMA) to fill in a createdmiddiaphyseal bony deficit and found torsional rigidities to be approximately 39% of control fora Kuntscher nail embedded in PMMA and approximately 18% of control for multiple Steinmannpins embedded in PMMA.

Obviously, a single IM Steinmann pin in an unreamed femoral medullary canal devoid of PMMA would have even less contact and mechanical interlock than the Kuntscher nail orstacked pins and would likely yield very low torsional resistance, depending of course on thefracture type. The torque capabilities of multiple (stacked) IM Steinmann pins conceivably couldapproach but not surpass the performance of the Kuntscher nail for transverse long-bonefractures. With an understanding of these mechanical limitations, it is clear that to optimizefracture stability using IM fixation with Steinmann pins, ancillary fixation may be indicated toprovide rotational stability. In the case of a closed pinning with a single IM Steinmann pin, thisnecessitates using an external splint or cast, which of course in combination may predispose to"fracture disease." For open single IM pinning the addition of one or more hemicerclage wirespotentially can afford adequate torsional resistance. Similarly, for multiple stacked pinning of oblique fractures, full cerclage wires provide essential rotational support. Because of therelatively low torsional resistance of IM fixation even under conditions of optimum wire and pinor nail placement, it is advisable to enforce strict confinement of the orthopaedic patient,particularly in the immediate postoperative period and for one month following or until thereexists radiographic evidence of healing.



















By virtue of location axially in the marrow cavity of long bones, IM fixation is excellently suitedto resist bending forces imposed on the bone and by this means maintains axial alignment. AnIM device, however, has no capacity to sustain axial compressive loads, and therefore themethod should be reserved for those fractures that can be reconstructed to transmit axial forcesthrough the surrounding bone, that is, good end-to-end contact.

As might be expected, the two major potential complications to the use of IM fixation arerotational instability and fracture shortening. Moreover, resultant motion at the fracture sitepredisposes to pin migration and may ultimately lead to nonunion or malunion. Additionally,should the migrating pin penetrate the skin, the fracture site may become infected, therebyintroducing the serious complication of osteomyelitis.

COMPRESSED PLATE FIXATIONThe need for early postoperative weight bearing in treating fractures of all types in animals hasled the veterinary surgeon to pursue fixation techniques that afford maximum postsurgical

rigidity. In the 1960s Lindahls(54-56) investigated the relative stabilities of various forms of internal fixation applied to oblique and transverse fractures of long bone. His conclusions, basedon results from flexural and torsional testing, were the following: no form of internal fixationcould match the flexural and torsional strength of intact bone; and of the methods available, plateand screw fixation provides the maximum rigidity.

Hayes and Perren(42) advanced the concept of "interfragmentary compression" to increase therigidity of the composite plate-bone system. By torsional testing of sheep tibia before and afterosteotomy with compression-plate fixation, they established a torsional rigidity of the plate bonecomposite approximately 60% of the control intact tibia and significantly greater than the plate-bone composite without interfragmentary compression. Results also showed that torsional

rigidities were optimized by placing screws close to the fracture or osteotomy site so as toincrease the length over which the bone and plate act as a composite system.

Similar studies(40,71) were extended to quantitation of the flexural rigidity of the plate-bonecomposite system. Four-point bending was performed on osteotomized and plated humancadaver femora and tibias, with and without interfragmentary compression, and results werecom- pared with the intact control. Bending forces were applied in two distinct configurations,one that tended to open the osteotomy line opposite the applied plate and one that tended to closethe osteotomy site. Results from a typical test specimen appear in Figure 12-29, in whichbending force is plotted against midspan deflection. Clearly, the plate-bone composite does notduplicate the flexural rigidity (stiffness) of the original intact tibia, confirming the results of Lindahl in an earlier study.(54) In the bending mode, which tends to open the osteotomy site, theresults indicate a greater stiffness of the compressed fracture than the noncompressed fracture inthe early phase of deflection. With increasing deflection, as bone-to-bone contact is lost, bothcompressed and non-compressed tibias demonstrate similar stiffness (slopes) corresponding tothe bending resistance of the plate alone. In the opposite mode, with bending tending to close theosteotomy site, the plate and bone as a composite become mechanically operational having anarea moment of inertia approximately equal to intact bone. Not surprisingly then, the compressedplate-bone composite has mechanical behavior and stiffness approaching that of the intact tibia.


















The noncompressed fracture experiences an initial low stiffness corresponding to plate bendingalone, followed by a slope similar to that of the intact tibia at deflections beyond that necessaryto close the osteotomy line. To minimize or eliminate this low stiffness portion of the curve, theauthors, following Bagby and Jones,(6) advanced the concept of "prebending or overbending"of the stainless steel plate in a concave configuration just over the osteotomy line. This

technique, upon screw tightening and subsequent compression, results in maximum contact of the cortical bone diametrically opposite the plate (a preload), producing increased flexuralrigidity. Aeberhard' extended these studies to torsional loading and demonstrated a markedimprovement in torsional rigidity using compression with prebending of the plate.

FIG. 12-29 Plot of bending force versus midspan deflectionfor compressed and noncompressed osteotomies relative tointact bone. Results suggest that plates should be applied insuch a way that the forces acting on the bone tend to close thefracture site. (Hayes WC: Biomechanics of fracture treatment.In Heppenstall B (ed): Fracture Treatment and Healing, p 124.

Philadelphia, WB Saunders 1980)

From these studies the authors concluded that for maximum mechanical stability of a transversefracture, bone contact is optimized by prebending the plate prior to application. In addition, sincethe bone-plate composite showed greater flexural rigidity in the "bending closed" loading mode,it was recommended that these devices be applied such that the acting forces in the system tendto close the fracture or osteotomy site. Usually this necessitates applying the plate on the convexor "tension band" surface of the bone. These tension band surfaces have been determined to areasonable degree of accuracy by direct measurements in vitro and in vivo(47,48,92)and bytheoretic mechanical analyses.(24,30,62,71) Along the same vein as Hayes and Perren, Bynumand co-workers(17) investigated the biomechanical capacity of installed commercial bone

fixation plates applied to the dorsal surface of equine third metacarpal bones. Bone specimens,having midshaft osteotomies and compression plating, were loaded to failure in compression,torsion, and flexure. Flexural loads were applied in the bending open and bending closed modesas well as from the side, or lateral loading. Results of testing, although lacking statisticalanalysis, demonstrated a range of plate-bone composite strengths from 16% to 67% of wholeintact bone. As expected, the lowest strength corresponded to the bending open loading modeand the highest, to the bending closed mode. The authors concluded that plate- bone compositestrength is dependent on the type of fracture and the mode of loading and that state-of-the- artplate fixation per se lacks the necessary mechanical requirements for successful equine fracturestabilization. Obviously, in long-bone fractures of small animals and humans in whomtransmitted loads are significantly smaller and for whom confined rest or restricted activity is

more simply enforced, the margin of safety in the use of plate and screw fixation is markedlyimproved.

The same authors,(76) in a subsequent investigation, performed parametric studies with varyingplate length, plate width, and screw diameter, again using the osteotomized equine metacarpusbut in flexural loading only. Results suggested that with optimized experimental parameters of plate length, width, and screw diameter the strength of the reconstructed osteotomy does notexceed 60% of the intact bone. Additionally, it was demonstrated that plate width (from 0.5 inch-






















1 inch) and screw diameter (from No. 8 to No. 14) did not significantly influence the strength of the plate-bone system. Increasing plate length, however, from 3 inches to 6 inches improved theflexural strength of the fixation approximately twofold. The finding emphasizes the need toselect plate length such that bending forces can be more uniformly distributed along the length of the plate-bone composite.

The question of optimum plate placement for long-bone fractures has long been a topic of discussion among orthopaedists. Hayes and Perren(4l,7l) suggested positioning plates on theconvex or tension band surface of long- bone fractures to provide maximum resistance tobending of the plate-bone composite. Minns and co-workers(61) pursued this concept further,hypothesizing that plate placement on the tension surface (of the tibia) is mechanically morestable in both flexural and torsional testing than placement on the compression surface. Theoreticcalculations drawn from gait analysis data by Paul(68) plus inherent geometric properties of human tibia supported the contention that the anterolateral surface of the tibia experiences themaximum tensile stresses during gait. Subsequent studies to test their hypothesis and the validityof their theoretic calculations were performed using physiologic loads and various plate

configurations. Results indeed confirmed that plate placement on the anterolateral surface of thetibia provided more resistance to physiologic load than on the anteromedial surface. In addition,it was concluded that tibial osteotomies fixed with compression plating were more rigidlystabilized than those without compression. Further, in comparing plate-bone rigidity in relationto plate size, it was found, not unexpectedly, that a thicker plate resists bending and torsion betterthan a thinner plate.

Whether these biomechanical findings in human tibias find practical application in veterinaryorthopaedics depends on factors such as the increased complexity of plate application on theanterolateral side of the tibia and the extent of improved clinical performance of this platingconfiguration. After discussion with many veterinary surgeons in clinical practice, it is apparentthat the large majority of clinicians do not feel justified in abandoning the anteromedialplacement of plate fixation of tibial fractures.

MATHEMATICAL MODELINGThe high expenditure of time and money in conducting scientific biomechanical experimentationhas led investigators to pursue theoretic mathematical models that can be fitted to the variousmechanical and material conditions of internal fixation. Very simply, the technique utilizesavailable information from experimentation pertaining to internal fixation combined withmathematics of theoretic mechanics to approximate or predict the biomechanics of internalfixation relative to various specific parameters. The most popular method to this end has been inuse since the early 1970s and is termed the "finite" element method. The following is a brief review of information derived from this method relevant to fracture fixation biomechanics.

Finite element modeling of transversely osteotomized and compression-plated equine thirdmetacarpals led to the early realization that axial loads, mathematically applied, had predictableresults on the plate-bone composite; namely, tensile forces tend to open the osteotomy line whilecompressive forces tend to close the osteotomy site.(78)These results supported the purelyexperimental findings cited previously.(17,24,26,71) The analysis also predicted that under




















conditions of axial compressive loads the plate carries approximately 18% of the load, therebyprotecting the underlying bone from experiencing the total force transmitted.

In similar modeling of an oblique fracture of the equine third metacarpus, the stress distributionof the contacting fracture ends was investigated as a function of plate tension, bone screws, and

axial forces.(79) Application of axial compressive force was found to increase the area of contactas well as the magnitude of contact stresses at the osteotomy site. Plate tension alone, however,in the absence of axial compressive loads on the bone resulted in very high stress concentrationsbeneath the plate with distraction and loss of contact of cortical bone opposite the plate. Theinsertion of a single interfragmentary screw perpendicular to the bone axis tended to antagonizethese unwanted effects of compression plating.

Askew and co-workers(5) used photoelastic and mathematical analyses to quantitate thephenomenon of bending caused by the eccentric placement of a compression plate on a longbone. In the case of a straight plate applied to a straight bone subjected to physiologic axialbending loads, eccentric compression plating resulted in extremely high stress concentrations

beneath the plate in the order of 107 N/m2 and a decrease in bone cross- sectional contact to 8%to 20% of normal. Based on their analysis the authors supported the prebending of compressionplates prior to application as advanced by Perren and Hayes.(4l) Plant and Bartel mathematicallyevaluated the bending effect of compression plating as a function of plate modulus.(74) Resultsof analyses of metallic plate-bone composites corroborated the above conclusions of Askew andco-workers: high stress concentration beneath the plate and fragment distraction of the "trans"cortex. Moreover, with plate modulus matching that of bone, compression was found to furtherexaggerate the bending effect, producing even larger osteotomy distraction than observed withmetallic plates. These results were further substantiated by Hayes and co-workers, who showedthrough an idealized bone-plate model that overbending plates even minimally (approximately0.2¡) produces more nearly uniform compressive contact stresses across the osteotomysite(4l)(Fig. 12-30). Clearly, then, experimental and mathematical evidence strongly supports theclinical practice of overbending compression plates for application on transverse or short obliquelong-bone fractures. The resultant plate-bone composite demonstrates increased bone contact atthe fracture site, more uniform end-to-end stress distribution, and less cyclic bending of themetallic plate.

PLATE-INDUCED OSTEOPENIAAs mentioned in the introduction to this chapter, compression-plate fixation shares with the bonethe job of force transmission across the fracture line. This phenomenon, so-called stressprotection of bone, results from the large mismatch of modulus between bone and metallic plate.Because mechanical stress is integral to bone remodeling and form (Wolff's law), a shielding of the bone from stress is postulated to cause a phenomenon termed plate-induced osteopenia. Thistopic has been the subject of several investigations, theoretic (mathematical) as well asexperimental.(2,26,76,91,93,99)




























FIG.12-30 Contact stresses at an osteotomy site for straight andprebent plates. For straight plates, the stresses are highlycompressive directly beneath the plate. Prebending the plate resultsin a more nearly uniform distribution of contact stresses. ( HayesWC: Biomechanics of fracture treatment. In Heppenstall B (ed):

Fracture Treatment and Healing, p 124. Philadelphia, WB Saunders.1980)

Woo and co-workers(103) used finite element analysis to evaluate the effect of plate modulus onlong-bone remodeling. Earlier quantitative histologic studies(2) demonstrated gradations of plate-induced osteopenia in the lateral cortex of canine femora subjected to lateral compression-plate fixation. The degree of osteopenia was postulated to correspond to the modulus of thecompression plate applied such that high modulus, for example conventional stainless steel,produces marked cortical thinning while lower modulus plates are associated with lessosteopenia. Subsequent finite element analysis modeling of the human femur has yielded thetheoretic magnitude of stress protection directly beneath the plate. In a comparison of bone stress

beneath plates having elastic moduli differing by an order of magnitude (vitallium versusgraphite-fiber-reinforced composites) it was determined that the vitallium plated boneexperiences only 7% of the stress of weight bearing when compared with the control unplatedfemur while composite-plated bone experiences 53% of weight bearing stress. The authors usedthis calculated stress reduction to explain the observed relative osteopenias in the canine femoraof the previous study.

These theoretic calculations were confirmed experimentally by Cochran(76) and by Schatzkerand co-workers(87) using in vitro strain gauge measurements of plated canine femora subjectedto axial loads. Cochran demonstrated an 84% reduction in bone strain immediately beneath ananterolaterally placed, four-hole AO plate and a 22% reduction on the medial side. Schatzker and

colleagues evaluated bone strain as a function of plate type: four-hole semitubular plate versusfour-hole AO dynamic compression plate. Strain measurements beneath the semitubular platewere reduced 64% on the laterally plated surface and 20% on the medial cortex relative to themeasured strains in intact bone. The dynamic compression plate, on the other hand, reduced bonestrain even more-77% in the lateral cortex and 34% on the medial cortex.

The results, both theoretic and experimental, document the potential of internal fixation viaplating to protect the underlying bone from normal physiologic stresses of muscle activity andweight bearing. The absence of sufficient mechanical stimuli may perhaps explain the osteopeniaobserved by Akeson and co-workers(2) beneath the plate in canine femora. A plausiblealternative explanation, however, may be that the plate acts as an impediment to vascular supply

to the bone beneath the plate or simply as a foreign body producing secondary or chemicalosteopenia unrelated to mechanical stimuli. Obviously the possibilities require furtherinvestigation.

In summary, internal fixation using plates and screws yields the most stable form of fixationimmediately post-operatively,(50,59) although no means of fixation is mechanically capable of achieving the strength and stiffness of whole intact bone. Whether or not this rigid fixation is
























detrimental to the subsequent development of late-stage fracture strength will be discussed in asection to follow.

BIOMECHANICS OF SURGICAL SCREWS

Surgical screws in veterinary orthopaedics are commonly employed in fracture management tosecure internal fixation plates to bone or to appose bony fragments pursuant to the principles of interfragmentary compression. Less commonly, screws may be used as anchoring points toattach synthetic ligaments or tendon for joint reconstruction. The biomechanics of surgicalscrews have received considerable attention in the literature, yet owing to the many variablesinvolved in determining optimum screw design, the present armamentarium of commerciallyavailable screws has resulted as much from empirical trial and error as pure scientific endeavor.Factors to consider relative to strength and holding power in bone include screw diameter (bothcore and outside diameter), thread type (angle, cross-sectional configuration), technique of application (pretapped versus self-tapping, pilot hole size) and of course inherent bonecharacteristics (thickness of cortex, bone damage upon screw introduction, and cortical versus

cancellous holding strength at the time of screw introduction and later in the healing phase).Optimization of all parameters clearly would necessitate a parametric study of enormouscomplexity. Fortunately in clinical veterinary orthopaedic practice, assuming the tenets of internal fixation are closely observed, failures of internal fixation, when they occur, rarely aredue to screw deficiencies either in inherent structural strength or holding force in bone.

Surgical screws are used to compress plate to bone or bone to bone and therefore by designexperience large tensile loads. To achieve optimum tension, however, a screw during the processof insertion undergoes a complex set of forces, including axial compression to maintainmechanical interlock between screwdriver and screw and torsion to drive the screw and manifestcompression of components. Torsional force can be specifically subdivided into the forcenecessary to cut threads (in the case of self-tapping screws); to overcome thread friction; toovercome friction at the screwhead/countersink interface; and ultimately to achieve the desiredtensile force in the screw(40,43) (holding power in bone). Hughes and Jordan have shown that aslittle as 5% of the total applied torque upon insertion is converted into screw tension under theadverse conditions of no pretapping of the pilot hole and no lubrication. To optimize theefficiency of torque conversion to screw tension, the authors suggested using the largestpracticable pilot hole, pretapping the screw hole, and irrigating with saline as a lubricant.(43) Using these measures a 65% conversion of applied torque to useful screw tension can beachieved. Furthermore, they suggested, from the results of their testing of screw insertion forces,that the maximum torque necessary to drive a screw should not exceed 65% of the screw'sultimate strength. This figure was determined after appreciahng that screw failure under theinfluence of multiple loading modes, such as torsion plus tension, may occur at torsional loadswell below the screw's ultimate strength in pure torsion.

The orthopaedist must acquire the surgical judgment to select the proper screw size based on themechanical demands placed on the screw in performing its intended function and the size andquality of the bone into which the screw is to be inserted. Hughes and Jordan have shown thatscrew holding power is dependent on the shear strength of the surrounding bone and isindependent of the mechanical properties of the screw.(43) Not surprisingly, cortical bone














demonstrates the highest holding power,(17) reported by Gotzen to be approximately 400 N/mmbone thickness for the AO 4.5-mm screw.(39) Cancellous bone is much weaker, yielding aholding power of 4 N/mm7 of surface.(70) Accordingly, to optimize pullout strength incancellous bone it is recommended that screws have a wide, deep thread and high threadpitch.(70) A useful concept in describing screw-to-bone contact is the parameter termed

interference. Interference is defined as the difference between the major diameter of the screwand the pilot hole diameter divided by the major diameter minus minor (core) diameter. Forexample, for the AO 4.5-mm cortical screw,

Interference = (4.5-3.2)/ (4.5-3.0) x 100 = 87%

For practical purposes, interference can be thought of as the degree of interdigitation of screwthread with bone. Interestingly, studies to evaluate pullout strength in cortical bone as a functionof interference have shown a poor correlation,(17,26,43,47) confirming the work of Hughes andJordan(43) that holding power is proportional to the surface area of cortical bone subjected toshear stress and therefore the major diameter of the screw. Screws having small core diameter

and high interference, although appropriate for cancellous bone, are subject to torsional failureon insertion into dense bone as a result of high torsional forces due to increased thread-bonefriction. Bynum and co-workers(17) have suggested interference values of 35% to 50% foroptimum pullout strength of pretapped machine-type screws in cortical bone. Nunamaker andPerren evaluated screw holding power in bovine cancellous bone and concluded that screw corediameter should be as large as conditions allow for maximum screw strength and that the majorscrew diameter should be maximized to achieve the optimum holding strength.(65)

Recognizing that screw pullout strength is a function of the shear properties of the surroundingbone and that these properties may potentially change with time from the trauma of insertion orwith healing, Schatzker and co-workers(86) conducted an in vivo study of screw holding power

in mongrel dogs for periods of up to 12 weeks. At no time did holding power drop below thetime "0" value, and at 6 and 12 weeks, pushout strength exceeded the initial postoperative values.Not unexpectedly, the 4.5- mm AO screw, having the largest major diameter in the series of screws tested, exhibited the greatest holding power. Subsequent histologic evaluation of thethread-bone interface revealed no impairment of the bone-healing process although lacunaewithin 1 mm of the screw threads were found to be devoid of nuclear material in the early (6weeks) healing phase. Although results of this study must be tempered by the fact that the screwswere only minimally loaded during the experiment, it is apparent that the trauma of the screwinsertion was not deleterious to holding power over a 12-week period.

In an investigation of screw tension associated with plate fixation, Laurence and co-workers(50) reported on engineering considerations related to the internal fixation of fractures of the tibialshaft in humans. Osteotomized human tibias were fixed with plates and specially instrumentedscrews and taken to failure in a bending "open mode." Results indicated that the maximumtension generated in the screws was only about one half the potential pullout strength of thescrew and much less than the ultimate strength of the screw. The authors concluded that fourscrews (eight cortices) were sufficient to provide adequate stability to a plated transverse tibialfracture at the moment of implantation but that more screws may be necessary to ensureoptimum stability over time. This work supports the empirical impression in veterinary










































orthopaedic practice that inherent screw deficiencies are rarely responsible for the failure of internal fixation of fractures.

BIOMECHANICS OF FRACTURE HEALINGThe biomechanics of fracture-fixation methods have been discussed with particular emphasis onimmediate post-fixation stability. Although of extreme value, this information requiresinterpretation in light of the biomechanics of the subsequent fracture healing process. That is, therate (indeed ultimate extent) at which a fractured bone unites is determined in large part bymechanical factors in and near the fracture environment. The process of fracture healing hasbeen thoroughly covered in a preceding chapter. (See Chapter 3.) The orthopaedist'sresponsibility is to reconstruct the normal anatomy of a fractured bone and to provide at least theminimum stability necessary for healing to proceed. This entails limiting fracture fragmentmotion sufficiently to maintain the viability of the interposed tissue. The presence or absence of motion determines the type of tissue that can survive between fracture ends, which in turn

dictates the type of bony union to be expected: primary, secondary, or nonunion. Whether or nota tissue type survives in the face of motion depends on its inherent ability to withstand strainsboth normal and angular. Perren has reported(69,70) that granulation tissue, cartilage, and bonecan withstand normal strains of 100%, 10%, and 2%, respectively, and angular strains of 40¡, 5¡,and 0.5¡, respectively. In a poorly immobilized fracture in which large tissue strains are to beexpected, the mechanical properties of granulation tissue make it more suited to survive, thusconstituting a nonunion. In the sequence of endochondral ossification leading to secondary bonyunion, a continuum of increasing stiffness (low strain tolerance) and strength is observed intissue types from granulation tissue to cartilage and finally to bone. Thus, "the precursor tissuesprepare the fracture gap mechanically and biologically for solid bone union."(70) In the case of rigid internal fixation in which deformations between fracture ends are limited to less than 2%strains, bone can form directly along vascular elements to a so-called primary bone union withno interposed fibrous granulation or cartilaginous phases. This phenomenon has beendemonstrated by Perren and co-workers in an evaluation of cortical bone healing in sheep usinginstrumented compression plates.(72) Histologic results as shown in Figure 3-8 reflect themechanical stability at the osteotomy site. Remodeling osteons are seen directly traversing thefracture line without the presence of fibrous or cartilaginous components.

Traditionally the clinical progress of fracture healing has been documented by radiographicevaluation and clinical empiricism. Experimentally the stages of fracture healing have beenclassified based on biochemical observations, histologic appearances, and radiographic criteria.White and co-workers(94) were the first investigators to correlate histologic and radiographicstages of fracture healing with objective measurements of the mechanical properties of healingbone tested to failure. The methodology utilized external skeletal fixation and was originallydeveloped in an attempt to quantitate the effect of cyclic loading on the rate of fracture healing.To this end the investigation fell short of its mark. Nevertheless, results of torque-anglemeasurements and radiographs made after mechanical testing of tibial osteotomy sites in rabbitsafforded a clear delineation of four biomechanical stages of fracture repair. Figure 12-31 is acomposite torque-angular displacement curve of six representative rabbit tibias at variousfracture healing times. Stage I repair corresponds to short fracture healing times (< 26 days) in






















which ultimate loading of partially healed bones results in failure through the originalexperimental fracture site with a low-stiffness, rubbery pattern. Stage II healing is marked byfailure through the osteotomy site as in Stage I but with progressive development of a high-stiffness, hard-tissue pattern as shown at 27 days in Figure 12-31. Stage III repairs undergotorsional failures partially through the original experimental fracture site and partially through

the previously intact bone with a high-stiffness, hard-tissue pattern (49 days, Fig. 12-31). StageIV fracture healing again displays a high-stiffness pattern upon mechanical testing (56 days, Fig.12-31); however, failure occurs radiographically at sites unrelated to the original osteotomy site.

FIG. 12-31 A composite torque-angle graph of six bonesrepresentative of the entire bone-healing period. The numberson the graph indicate days of healing time. As healingprogresses, there is increase in the strength of union as shownby the changes in the torque-angle graphs. (White AA, PanjabiMM, Southwick WD: The four biomechanical stages of fracture repair. J Bone Joint Surg 59A:188, 1977)

Statistical analysis of data demonstrated the four stages to correlate closely with fracture strengthand healing time. The authors emphasized the need for a nondestructive means of assessing theclinical progression of fracture healing. For the present, the veterinary orthopaedist must rely onradiography, clinical judgment, and the empirical passage of time to monitor fracture healing. Onthe horizon, however, are promising nondestructive, fracture strength assessment techniquesincluding ultrasound, stress wave propagation, and resonant vibration.(49)

In a later study White and co-workers(95) used a similar rabbit model to investigate temporalchanges in physical properties of healing fractures. Pertinent results indicate that the maximumtorque to failure at 3 weeks is 25% of intact bone, progressing to 75% at 9 weeks. Angular

deformations prior to failure are understandably very large at 3 weeks and diminish to slightlyless than intact bone at 9 weeks. This finding can likely be explained by the large polar momentof inertia associated with callus formation.

It should be recognized that the testing methodology employed in this series of investigationsinvolved the use of a transfixation apparatus and various fixed and cyclical compressive loads.Also, rabbits were not permitted normal weight bearing on their hind limbs. Whether theseresults then can be extrapolated to a comparative assessment of clinical fracture healing usingstandard internal, external, or transfixation is open to question. Minimally, the significance of these reports would support the continued use of radiography as an indicator, albeit crude, of fracture strength and fracture healing progress.

Along a similar vein, Whiteside and co-workers(98) investigated the biochemical characteristicsof fracture callus temporally and correlated these findings with biomechanical data and theradiographic stage of fracture healing. Specifically, osteotomized rabbit tibias repaired with IMpins were subjected to tensile testing at various stages of fracture repair (determined temporallyand radiographically), and results were correlated with the biochemical composition of callus. Asmight be expected, in the early stages of callus formation the observed large total strain to failurewas positively correlated with high mucopolysaccharide level (hexosamine) and low calcium














level. As callus matures, however, tensile strength increases progressively and the callusbecomes stiffer and therefore can endure less strain to failure and correspondingly less energy tofailure. Associated with these changes the biochemical picture is one of diminishing hexosaminelevel (particularly in the callus-bridging phase) and increasing calcium content of the callus.Interestingly, in the final stage of callus maturation no biochemical changes were observed, yet

mechanical properties changed markedly toward increased strength and stiffness. It is likely thatthis phenomenon is a result of the internal remodeling process that works to optimize thestructural characteristics of the healing fracture.

Additional and often underestimated factors in determining the rate of fracture callus formationand maturation are the extent of soft tissue trauma in the vicinity of the fracture and the type of surgical approach used to achieve fracture reduction and stabilization. For example, soft tissuetrauma and extraperiosteal dissection have been shown to retard the formation and strength of fracture callus when compared with control fractures approached subperiosteally.(97) Clearly itis important to be aware of such an influence when designing or interpreting comparativeexperiments of fracture healing.

THE ROLE OF FUNCTIONAL WEIGHT BEARING

It has long been theorized that fracture healing rate and ultimate fracture strength may beprofoundly influenced by mechanical factors in the fracture environment. In recent years severalexperimental investigations have been conducted to elucidate these mechanical factors with theintent of optimizing the fracture healing process(32,44,51-53,66,67,75,91,95,101) White and co-workers, using a rabbit osteotomy model previously described,(67) attempted to demonstratesignificant mechanical differences between fractures healed under conditions of constantcompression and fractures subjected to constant compression and a superimposed cyclic load.

Constant and dynamic loading was implemented by means of a specially designed externalfixator. Previous investigations had been unsuccessful in demonstrating a significant effect.(95) In their most recent study, by imposing more dissimilar mechanical environments, they (99) were able to demonstrate at 6 weeks of healing significantly increased torque and energyabsorption to failure (as well as lower stiffness) in fractures exposed to cyclic loading versuspairmates treated with constant compression alone. No significant differences, however, werenoted at 4 or 8 weeks. There was a "suggestion" that compression-treated bones may be strongerthan load-cycled bones in the earlier phases of fracture healing. This finding had also beenreported in an earlier study by Sarmiento and co-workers.(85) Apparently, in the early phase of fracture healing, cyclic loads and the resultant tissue strains retard the normal and delicateprocess of vascular repair and granulation tissue formation,(27) thus delaying creation of an earlyfibrous callus. However, in the interval between 4 and 6 weeks cyclically loaded fracturesdemonstrate superior mechanical behavior in both torsional testing(72,99) and three-pointbending.(85) Again, this is probably due to the more extensive strain-induced callus andattendant increased inertial properties. The positive correlation between motion and increasedfracture callus and cartilage is compatible with the findings of a number of previous noteworthyinvestigations.(36,57,60,104) In 8 weeks, however, the effect disappears.











































It is conceivable from these studies that the optimum fracture-fixation method would incorporaterigid immobilization in the early revascularization stage of fracture healing followed at 2 to 6weeks by the introduction of small interfragmentary strains (via controlled weight bearing) tomaximize callus formation and thereby speed the return of strength and stiffness. At 8 weeks andbeyond it appears that the process of fracture reorganization and remodeling leads to the same

mechanical end point regardless of function and immobilization history. Cyclic loading,therefore, may be useful in hastening the return of fracture strength add stiffness but does nothave a significant effect on ultimate fracture strength of the remodeled fracture site. It remainsunclear whether this accelerating effect is attributable to motion at the fracture site, the variableapplied force, or perhaps a strain-induced biologic signal (e.g., bioelectrical).

Additional information on the influence of motion on the formation of fracture callus and theresultant rate and strength of fracture healing was presented by Piekarski and co-workers.(73) Their original objective was to test biomechanically the effects of delayed IM internal fixationversus immediate internal fixation. In evaluating results from the delayed fixation of osteotomiesof the radius of rabbits, no significant mechanical difference was observed within the 6-week

experimental period between fractures fixed immediately and those fixed one week following acreated fracture. There was, however, a noticeable increase in callus cross section and callusvolume in the delayed-fixation group, particularly in the early stages (up to 3 weeks) of fracturehealing. Not surprisingly this cross section diminished in size with the added support of the IMdevice; however, this was not evident until 2 weeks following insertion. Although callus wasmore exuberant in the delayed-fixation group, its material strength was less than that of thecontrol (immediately fixed) group at 3 weeks as a result of the discovery of an increased callusporosity. This porosity was found to diminish slowly from 3 to 6 weeks, and in conjunction withthe observed reduction in cross section in this period the delayed-fixation group demonstratedreduced failure loads and tissue strengths. Therefore, although delayed fixation appears toenhance early callus formation and thereby increase fracture stability, the poor quality of callusand the rapid reduction in cross section over time would suggest that its use has no particularadvantage over immediate fracture fixation.

In correlating mechanical behavior with radiographic fracture appearance, Sarmiento and co-workers(85) found that the radiographic disappearance of the fracture line in the rat tibialosteotomy model did not necessarily indicate completed fracture healing based on biomechanicalcriteria. That is, tibias subjected to functional weight bearing in which a radiographic fractureline was visible were mechanically stronger than immobilized rat tibias in which the fracture linehad disappeared. Again this is undoubtedly due to the superior structural properties in theexuberant noncalcified fracture callus. The finding, however, would warrant caution in assessingfracture strength solely on the basis of the presence or absence of a fracture line.

In summary, biomechanical information is accumulating to support the functional loading of fractures, except perhaps in the immediate postfracture phase. Fractures treated with functionalloading demonstrate a more rapid return of strength and stiffness when compared withimmobilized fractures, although the biomechanical end point to fracture healing (afterremodeling has occurred) is the same regardless of fixation technique. An additional advantageof functional weight bearing, of course, is the maintenance of normal muscle mass and freedomof joint motion. Weight bearing and the consequent fracture motion appear to correlate positively











with the size of the fracture callus. Some investigators suggest that the continued mechanicalstimulation of weight bearing accelerates the subsequent callus ossification process; however, thevalidity of this impression awaits rigorous scientific proof.

These theoretic and experimental findings are perhaps not too surprising to the veterinary

orthopaedist who historically has been managing fractures in animals in the face of anoverwhelming desire on the part of the animal to ambulate. Functional weight bearing is andalways has been recognized, at least empirically, as an integral part of successful fracturemanagement in animals. Studies that advocate limited though appreciable motion at the fracturesite to enhance callus formation would be difficult if not impossible to implement in animalsowing to lack of patient cooperation. Although theoretically, functional weight bearing results inan earlier return of strength and stiffness in a controlled situation in experimental animals, inclinical animals there is no such control. Accordingly the veterinary practitioner must tailor thefixation method to the type of fracture and the type, size, and expected activity level of theanimal. Although rigid internal fixation with plate and screws may inherently antagonize thedesirable callus formation process, it nevertheless is clearly the treatment of choice in the

extremely active animal whose physical activity would exceed the mechanical constraints of anyalternate form of fracture treatment. It must be remembered that in time the reorganization andremodeling process yields mechanically the same result regardless of function andimmobilization, and therefore the appropriate choice of fixation and activity level is that whichhas the capacity to provide adequate stability for the minimum duration from fractureorganization to callus reorganization.

INTERNAL FIXATIONThe mechanical influence of internal fracture fixation on the biologic healing process has beenthe subject of considerable investigation, yielding valuable information to the orthopaedicsurgeon. Fortunately for the veterinarian, most of these studies have been performed inexperimental animals and therefore the results (particularly in the canine) are directly applicableto veterinary orthopaedic practice.

PLATE AND SCREW FIXATIONIt has been stated previously that compression-plate fixation facilitates rigid internal stability,leading to primary bone union and absence of callus. What, however, are the biomechanicaleffects of compression on bone over time? Perren and co-workers,(72) using strain gauge,instrumented compression plates in sheep, characterized the time-dependent decay of interfragmentary compression forces postsurgically and evaluated subsequent radiographic andhistologic changes attributable to the compression. Strain gauge results indicated that post-surgical interfragmentary compressive forces of 60 kiloponds to 140 kiloponds diminished withtime but did not drop to zero over the 1 2-week experimental period. Interestingly, similar resultswere obtained when instrumented compressive plates were applied to intact (nonosteotomized)bone. This finding would suggest that the positive stabilizing effects of interfragmentarycompression are manifested, albeit in declining magnitude, throughout the critical first 3 monthsof healing. Also, histologic and radiographic results confirmed that in the force range of 60kiloponds to 140 kiloponds, no "pressure necrosis" of bone could be demonstrated at the fracturesite. This finding ran counter to many popular beliefs at the time (1969).







With the introduction of rigid plate fixation much practical concern has been voiced byorthopaedists regarding the phenomenon of stress protection and the attendant plate-inducedosteopenia. Particular concern stemmed from the practice at that time of double-plating femoralfractures, which resulted in an unacceptably high incidence of refracture following plateremoval. The topic of plate-induced osteopenia has been covered in a previous section on

mathematical modeling, however, pertinent experimental studies will be discussed below.

Investigations on plate-induced osteopenia have radiographically demonstrated the existence of cortical thinning beneath the plate for at least the first 6 months following application.(93) Histologic follow-up has pointed to a mechanism of increased periosteal resorption of corticalbone with associated increased bony porosity. These changes have been shown to be virtuallycompletely reversible within 3 months after plate removal.(93) Increased bone porosity was acommon finding in a study by Akeson and co-workers,(2) who demonstrated a relationshipbetween the degree of cortical porosity and the modulus of the applied plate (i.e., decreasingplate stiffness results in decreased bony porosity). Similar findings were reported by Tonino andassociates(91) in a comparison of stainless steel and plastic plates. Plastic-plated bones were

shown to have superior mineral mass and mechanical properties. On the other hand, stainlesssteelplated bones exhibited massive endosteal resorption as determined by microradiography. Incontrast to these findings, Woo and co-workers(102) were unable to demonstrate significantmechanical differences between 16 healed radial fractures treated with stainless steel plates andthose treated with low-modulus graphite-PMMA composites. A later study by this group(101), however, revealed a significantly increased amount of bony atrophy associated with healedfractures treated with stiff stainless steel plates in comparison with those treated with less stiff stainless steel plates. Interestingly, the material properties of healed fractures from bothexperimental groups were found to be indistinguishable. A detectable cortical thinning in thestiff-plated group was thought to be responsible for the significantly inferior structural propertiesobserved.

The effects of plating (with and without compression) on intact bone morphology wereinvestigated by Slatis and associates(90) using the rabbit tibia model. Histologically theapplication of a plate on an intact tibia was found with time to result in the formation of subperiosteal new bone and a concomitant resorption of subendosteal cortical bone. Pointcounting of cross-sectional areas of bone revealed no significant difference between compressedand noncompressed tibia by the end of the 9- months experiment. The net result, however, of plating intact bone regardless of compression was to increase the area of the medullary canal andat the same time to increase the total cross-sectional area of bone, creating a larger diameter,thinner walled tube.

Thus it appears that the application of a metallic plate to bone has the potential to producesignificant morphologic changes in the bone, including widening of endosteal haversian canals(increasing porosity), widening of the medullary canal, and enlargement of the overall outsidediameter. The precise mechanism or mechanical stimulus for these changes is unclear, but it isprobably related to the stress-protection phenomenon described previously. Of extremeimportance is the realization that similar morphologic alterations are to be anticipated along thelength of bone exposed to an overlying plate. This plate-associated weakening of bone warrantscareful discretion in removing metallic plates following radiographic evidence of fracture






























healing. Considering the weak correlation between radiographic assessment of fracture healingand the actual mechanical strength of the healed bone, optimum timing of plate removal is verydifficult to determine clinically. It has been shown in the dog that the mechanical andmorphologic changes associated with plate-induced osteopenia are completely reversible within3 months of plate removal.(63,93) Based on these findings, it is advisable to suggest several

months of restricted activity (if not temporary splinting) for the veterinary orthopaedic patientfollowing plate removal.

Contrary to the generally held views on plate-induced osteopenia, Noser and co-workers(64) have shown that plated osteotomized canine femora increase in torsional strength up to 6 monthspost implantation to a strength in excess of normal unoperated bone and thereafter decrease instrength apparently as a result of stress-protection phenomena. This would suggest that if peakmechanical maxima were identifiable clinically, the timing of plate removal could be determinedto minimize the risk of refracture. Presently such nondestructive noninvasive mechanical testingmethods are not available to the veterinary practitioner.

The existence of open screw holes following plate removal raises the question of the consequentmechanical weakening of bone. It has been shown that the mechanical influence of a freshlydrilled screw hole is relatively independent of size (if less than 30% of bone diameter) and for2.8- and 3.6-mm screw holes a 1.6-fold increase in stress is generated around a screw holerelative to the surrounding bone.(11) When freshly drilled bone is torsionally loaded, fracturetypically occurs through the drillhole and at approximately a 50% reduction in torque relative tothe contralateral control side(14) Burstein and associates(l4) histologically demonstrated densewoven bone filling screw holes in dogs within 4 weeks of screw removal, but despite this findingthey reported the radiographic existence of prominent radiolucent defects in bone even monthsafter screw removal. Apparently there exist significant density differences reflecting variation incalcium content between normal bone and the plug of bone filling a screw hole. After 4 weeksand the formation of a dense woven bone plug, however, torsional loading to failure producedfracture lines not localized to screw holes.

Conceivably the stress concentration effect of freshly drilled holes reported by Brooks(11) mayserve as a stimulus for adaptive bone changes in the area of the screw hole such that withhealing, stress concentration is eliminated. This impression was confirmed by a relatedexperiment in rabbit femora that indicated that bone adapts to reduce stress concentrationsregardless of whether the defect is an empty screw hole, a screw hole with screw in situ, or ascrew hole filled with silastic plug. By 8 weeks of healing, all femora so treated had similartorsional strength. If, however, after 8 weeks of healing with a metal screw in situ the screw isremoved, a significant (25%) decrease in mechanical properties is observed. Again, thisinformation would suggest that following plate and screw removal serious consideration shouldbe given to limiting the activity of the patient to allow the adaptive processes of the bone to reachcompletion.

IM ROD FIXATIONAs stated earlier, single IM rod fixation does not provide the rigid internal stability of plate andscrew fixation. Therefore more motion is to be expected at the fracture site, resulting in increased























callus formation and healing by secondary bony union. Curiously, very few investigations havebeen undertaken with the goal of optimizing the healing process relative to the many parametersof IM fixation. Of note is a study by Brown and Mayor, who investigated the biomechanicaleffect of IM rod fixation relative to rods of differing composition having vastly differentmoduli.(12) Rods were fashioned from stainless steel, titanium, Delrin, and nylon with respective

moduli relative to bone at 10-, 6-, 0.1- and 0.003-fold, respectively. Osteotomized tibias inrabbits were fixed using the above rods and allowed to heal for 16 weeks. Torsional testing withthe rods in situ demonstrated a 30% greater strength and 36% greater stiffness of fracturestreated with plastic rods versus those treated with metal rods. No mechanical differences werenoted between fractures fixed with the two types of metal rods nor between fractures fixed witheither type of polymer. From histologic and radiographic comparisons the authors correlatedthese mechanical differences with associated morphologic changes of the healed tibial fracture.Stiff metallic rods were found to elicit the formation of bony caps at the ends and a sleeve of bone along their length, while flexible polymeric rods caused no such changes. It was postulatedthat the metal rods, as a result of the formation of bony caps and sleeves, were capable of protecting the tibia from stress, thus explaining the observed lack of remodeling and loss of

cortical density relative to the polymer-treated fractures. Fifty percent of the plastic-treatedfractures achieved strengths within 10% of the control unoperated side within the 16-weekhealing period. Although rabbit tibial fractures without associated fibular fracture may representa rare clinical occurrence and may impart a significant component of rotational stability to thereconstructed fracture site, the results of this investigation would warrant continued study intothe feasibility of flexible IM devices.

Fracture healing associated with IM fixation has been shown to be influenced by the extent of soft tissue trauma as well as by the type of surgical approachÑsubperiosteal or extraperiosteal.Whiteside and Lesker investigated the rate of fracture healing and ultimate tensile strength of healed osteotomized rabbit tibia at 3 weeks following subperiosteal or extraperiosteal dissectionor complete muscle transection at the osteotomy site.(97) All osteotomized tibiae were internallyfixed using a single IM pin. Of the tibia exposed subperiosteally, 71% healed within 3 weeks,while of those exposed extraperiosteally only 7% had solid union. Muscle transection resulted ina marked retardation in fracture healing of all tibiae whether exposed subperiosteally orextraperiosteally. This finding, in combination with earlier work that demonstrated the adverseeffect of extraperiosteal dissection on the collateral circulation to traumatized muscle,(96) would strongly support a subperiosteal approach for both severely traumatized anduntraumatized extremities. Although not specifically studied, this conclusion would likely applyto all forms of internal fixation.

COMPARATIVE BIOMECHANICAL STUDIES:HEALING RELATIVE TO TREATMENT MODALITYFew investigations appear in the literature that attempt to quantitate mechanically the relativemerits of the various internal fracture-fixation techniques in terms of associated fracture healing.Furthermore, virtually no information in this regard is available on external means of fracturefixation. Clearly such information is of utmost importance to the orthopaedist who desires aknowledge of the relative mechanisms, strengths, and rates of healing to facilitate appropriateselection of fracture-fixation method.















Anderson,(4) in the mid-1960s, performed midfemoral osteotomies in dogs and histologicallyfollowed the fracture-healing process as a function of fixation technique. Five separate fracture-treatment groups were defined in increasing order of fracture stability: series 1, no fixation;series 2, loose-fitting medullary nails; series 3, tight-fitting medullary nails; series 4, plate andscrew fixation without compression; and series 5, compression plates. Results, as one might

expect, indicate that the mode of osteogenesis and the vascular and cellular response of healingwere largely dependent on the stability and reduction of the fracture fragments. Furthermore,eventual fracture-callus size was shown to be directly proportional to fracture instability. As anextension of this study, attempts at mechanical testing of fracture strength were performed in acrude fashion on plated femora with and without compression. A three-point bending load wasapplied gradually to healing excised femora by suspending a bucket from one end of the boneand slowly adding water. Although after 6 weeks the authors reported a union rate of 100% forthe compression-plated group versus 91% for the plated group having no compression, theycould not demonstrate a statistically significant difference in fracture strength using theirmechanical testing methodology. Nevertheless, from their results they suggested that thebeneficial influence of compression on fracture healing is to increase the rigidity of fixation by

impacting the bone ends and thereby reducing the interfragmentary space for bone bridging.A somewhat more sophisticated comparative study of fracture fixation relative to fracturehealing was performed by Braden and associates(9) also using the dog midshaft femoralosteotomy model. Fractured femora were treated with three different fixation methods appliedover three different intervals: IM pin in place for 6 weeks, IM pin with supplementary one half Kirschner apparatus removed at 4 and 6 weeks, respectively, and bone plates removed at 10weeks. All animals were killed at 10 weeks and the excised femora were tested in torsion tofailure to determine ultimate strength and stiffness (normalized to the contralateral unoperatedfemur). Results demonstrated the femora with IM pin and one half Kirschner apparatus to havethe highest strength (80.2% of the normal side) compared with IM pin alone and bone plate(61.9% and 36%, respectively). Stiffness results followed a similar pattern. The authorsconcluded that the two techniques of IM pin fixation and IM pin fixation with supplementary onehalf Kirschner apparatus are mechanically superior to bone plating for fracture fixation in theexperimental time period of 10 weeks. Furthermore, they went on to attribute these mechanicaldifferences to a relative lack of functional loading (stress protection) associated with femoralplating (see section on plate osteopenia, above).

Although these conclusions appear plausible based on the experimental results, the authors'interpretations are subject to criticism in several areas. First, for a strictly comparable study, thevarious types of fixation should have remained in situ for equivalent periods of time. Indeed,subsequent data reported by the same group'¡ indicated a rapid increase in strength and stiffnessof plated femora following removal of the device and resumption of weight bearing. It remainsunclear, then, whether the biomechanical differences reported here are solely dependent onfixation technique, on postfixation functional loading, or on a combination of the two. Second,no mention was made of callus type, size or geometry, raising the question of whether motion atthe fracture site and exuberant callus formation may have contributed more to the obviousmechanical differences than the postulated concept of stress protection. Third, the describedmechanical testing methodology adds numerous variables to the determination of ultimatetorsional strength and stiffness when compared with standardized torsional testing in the absence











of cycling.(15) Specimens were subjected to extremely complex, if not uncertain, stress fieldsand therefore endured disparate failure modes not easily normalized to the control femur nornecessarily comparable from one group with another.

The validity of these criticisms was borne out in a study by Mullis and co-workers(63) in a

quantitative and comparative analysis of fracture healing under the influence of compressionplating versus closed weight-bearing treatment. Results indicated that fractures treated by closedreduction experienced large periosteal callus formation and after 1 month regained superiorstrength and stiffness compared with 1-month, compression-plated fractures (immediately afterplate removal). This difference, however, was virtually eliminated at 3 months when the strengthof fractures previously compression- plated was shown to approach that of fractures uniting byperiosteal callus.

The influence of stress protection as postulated by Braden to explain the inferior torsionalproperties of plated femora at 10 weeks was modified subsequently by the same group of investigators in a continuation study of the effect of time on healing strength of plated bone.(64)

Indeed, in marked contrast to the earlier 10-week results, plate removal from osteotomizedfemora at 6 months was associated with torsional properties actually exceeding the strength andstiffness of the contralateral control femur. Curiously, this effect was reversed at 9 months whenhealed fractures immediately post-plate removal demonstrated significant reductions in torsionalstrength and stiffness of 34% and 18%, respectively. The authors attributed the improvedstrength at 6 months to selective osteonal activation stimulated by the fracture-healing process,while the 9-month reduction in mechanical properties was again related to stress protectionphenomena. No attempt, however, was made to explain the earlier 10-week strength data in lightof this new information. It was concluded that timing of plate removal should optimally coincidewith the observed mechanical maxima but that these findings are preliminary and clearly moreresearch is necessary. Perhaps the recognized weakening influence of screw holes on bonestrength represents a more plausible explanation for the 10-week strength data.(11)

In a well-controlled comparative study of various fracture-fixation techniques, Burstein andassociates measured the corresponding return of torsional strength and stiffness of osteotomizedfemora using standardized torsional testing methodology. (16) The primary purpose stated was todescribe and evaluate a new fixation technique of pericortical clamping designed to provideinitial fracture reduction and stability comparable to compression plating but without the needfor cortical screw holes and the associated cortical weakening.

Clamp fixation devices consisted of two interlocking halves made of cobalt-chromium alloy thatwere secured to the bone by a system of embedded spikes and maintained by two screws in thedevice itself. For comparative purposes, three other internal fixation devices were used in thisstudy: a standard four-hole rigid AO plate; a similar plate machined to yield a tenfold decrease inbending rigidity; and a Kuntscher IM rod appropriately sized to fit the unreduced femoral canal.Devices were applied to transverse femoral osteotomies in adult mongrel dogs and torsionallytested at specified time intervals of 0 (device in place), 10, 16, and 20 weeks (device removed).A corresponding group of dogs had devices applied similarly but without osteotomy. Half of these animals had their femora removed and tested at 10 weeks while the remainder had the






















devices removed at 10 weeks and their femora tested at 18 weeks. Results of mechanical testingwere expressed as percentages of the values from the contralateral control femur.

Figure 12-32 shows the rate of recovery of torsional strength, stiffness, and energy to failure of osteotomized femora from the four experimental groups. Similarly Figure 12-33 demonstrates

the recovery of torsional strength, stiffness, and energy absorption to failure at 18 weeks (afterremoval of the fixation devices at 10 weeks). Consistent with studies already discussed, thetorsional strength of freshly plated osteotomized femora reflects the stress concentration effect of the screw holes, producing a 50% reduction in strength.(14) In subsequent tests at 10, 16, and 20weeks immediately following plate removal, this effect persisted for both the rigid and flexibleplate and is in marked contrast to the results of Noser and associates(64), who, as discussed,found torsional strengths at 6 months to exceed the normal contralateral bone. From Figure 12-33 it is apparent that the cortical- weakening effect of screw holes was eliminated by the healingprocess within 8 weeks of plate and screw removal from unosteotomized femora. In addition,unosteotomized plated bones at 10 weeks showed a significant reduction in torsional strengthand energy to failure as a function of plate rigidity as previously shown by Woo and co- workers.

(101) However, the postulated stress protection effect between rigid and flexible plate in thiscase accounted for only 16% of the loss in strength, implying that the weakening effect of screwholes is the major controlling factor. In fact, throughout the experiment, the strength of bonefixed by internal fixation plates never varied significantly from the value predicted from aconsideration of the screw holes alone, suggesting that stress protection although real, may be anoverly emphasized concept.

FIG.12-32 Rate of recovery of torsional strength stiffness, andenergy to failure of osteotomized canine femora relative tofixation method In plot A, mean torsional strength wasdetermined at "O" time with fixation devices in place All other

data in A B, and C correspond to testing immediately followingplate removal. Note the negligible torsional strength at "O"time of intramedullary Kuntschner-nail fixation. (Data adaptedfrom Burstein AH, Moseley CF, Robinson RP et al:Comparative mechanical performance of a new clamp fracturefixation device. Personal Communication, 1983)

FIG.12-33 Rate of recovery of torsional strength stiffness, andenergy to failure of intact (nonosteotomized) canine femoraexposed to four different fixation devices for 10 weeks. Thecortical-weakening effect of screw holes at 10 weeks is almostcompletely eliminated 8 weeks following plate and screw

removal. (Data adapted from Burstein AH, Moseley CF,Robinson RP et al: Comparative mechanical performance of anew clamp fracture fixation device. Personal Communication,1983)

The clamp and Kuntscher nail, although demonstrating low initial mechanical properties, wereassociated with progressive improvement of torsional strength, stiffness, and energy absorptionof bone over the 20-week course of the experiment. Moreover, compared with plate fixation,


















these two cortical-sparing devices did not produce as marked a reduction in the mechanicalproperties of the fixed unosteotomized femora and therefore would be less susceptible torefracture upon device removal.

Of the fixation methods tested, the proposed clamp system clearly exhibited superior

performance. Disadvantages include the need for a wide variety of clamp sizes to ensureadequate fixation of all size dogs and the difficulty in fixing severely comminuted, nonunion, ormetaphyseal fractures. The obvious advantages are adequate, immediate fracture reduction andstability without the weakening effects of screw holes or disruption of the medullary bloodsupply. The information from this study is extremely helpful in evaluating the relative merits of various fracture fixation techniques, particularly for application in veterinary orthopaedicpractice. It is probably the only investigation to date having sufficient control to yieldstatistically significant differences among fixation methods.(40)





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Etiology of Fractures