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MECHANICAL EVALUATION OF PLATE WORKING LENGTH IN A CANINE FEMORAL MODEL

By

PEINI CHAO

A THESIS PRESENTED TO THE GRADUATE SCHOOL OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT

OF THE REQUIREMENTS FOR THE DEGREE OF MASTER OF SCIENCE

UNIVERSITY OF FLORIDA

2012

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© 2012 Peini Chao

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To my family.

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ACKNOWLEDGMENTS

First, I would like to thank all the people that have contributed to the work described

in this thesis.

Dr. Antonio Pozzi as my advisor and mentor has been an inalienable part of this

project. His perpetual energy and enthusiasm in clinics and research had motivated me

through this research journey. He gave me the confidence and all the support, enabled

me to get through every difficulty in front of me.

I would like to acknowledge my thesis committee members. The valuable advice

from Dr. Daniel Lewis, always pointed out the key points. Without his support I couldn't

have achieved a milestone. I gratefully thank Dr. Bryan Conard for providing me major

support, both in knowledge and in techniques, of engineering science. Dr. MaryBeth

Horodyski’s help in statistics was really precious in this thesis.

Additionally, I appreciate Dan Barousse and Debby Sundstrom for their important

experience sharing and all the assistance in every little detail in this study. Last but not

the least, this thesis would be incomplete without mentioning my parents: Shereen Lu

and Pan-Hwa Chao. Their unconditional support could only be cherished by providing

them a good research work. My big brother, Josh Chao has been another powerful

support during this whole time. Being my second family here, James Wang and

Ting-Bing Wu has not only guided me through many difficulties, but also replaced them

with fun and joy.

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TABLE OF CONTENTS page

ACKNOWLEDGMENTS .................................................................................................. 4

LIST OF FIGURES .......................................................................................................... 6

LIST OF ABBREVIATIONS ............................................................................................. 7

ABSTRACT ..................................................................................................................... 8

CHAPTER

1 INTRODUCTION .................................................................................................... 10

2 BACKGROUND ...................................................................................................... 12

Basic Mechanics of Materials ................................................................................. 12

Force, Deformation, Stress and Strain ............................................................. 12 Stiffness............................................................................................................ 13 Fatigue Failure ................................................................................................. 16

Applied Biomechanics............................................................................................. 17 Biomechanics of Fracture Healing .................................................................... 17

Bone Healing under Conditions of Absolute and Relative Stability ................... 20 Factors Affecting Stiffness of the Plate-Bone Construct ................................... 23

Choosing the type of plate: locking versus non-locking plates ................... 23 Choosing the length of the plate ................................................................ 26 Effect of the position of screw placement in the plate ................................ 27

3 EXPERIMENT ........................................................................................................ 35

Effect of Plate Working Length on Plate Stiffness and Cyclic Fatigue Life in a Cadaveric Femoral Gap Model Stabilized with a 12-hole 2.4 mm LCP ............... 35

Specimen Preparation ...................................................................................... 36 Mechanical Testing .......................................................................................... 38

Data and Statistical Analysis ............................................................................ 39 Results .................................................................................................................... 40

Discussion .............................................................................................................. 42

4 CONCLUSION ........................................................................................................ 48

LIST OF REFERENCES ............................................................................................... 54

BIOGRAPHICAL SKETCH ............................................................................................ 64

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LIST OF FIGURES

Figure page 2-1 Load-deformation curve ...................................................................................... 29

2-2 Load-deformation curves of three different material: bone, glass and metal. ..... 30

2-3 Area moment of inertia illustration ...................................................................... 31

2-4 Area moment of inertia between different implants ............................................. 32

2-5 Fatigue behavior in an S/N curve ....................................................................... 33

2-6 Absolute stability affects bone healing ................................................................ 34

2-7 Relative stability affects bone healing ................................................................. 34

3-1 Testing constructs of long and short plate working length stabilization techniques. ......................................................................................................... 46

3-2 Mean ± SD values of construct axial stiffness (N/mm) between long and short plate working length stabilization techniques ...................................................... 46

3-3 Photographs and free body diagrams of a long and short plate working length constructs before and after load to failure test .................................................... 47

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LIST OF ABBREVIATIONS

ANOVA ANALYSIS OF VARIANCE

AO-ASIF ARBEITSGEMEINS CHAFT FÜR OSTEOSYNTHESEFRAGEN – ASSOCIATION

FOR THE STUDY OF INTERNAL FIXATION

DCP DYNAMIC COMPRESSION PLATE

LCP LOCKING COMPRESSION PLATE

MTS MATERIAL TESTING MACHINE

SD STANDARD DEVIATION

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Abstract of Thesis Presented to the Graduate School of the University of Florida in Partial Fulfillment of the

Requirements for the Master of Science

MECHANICAL EVALUATION OF PLATE WORKING LENGTH IN A CANINE FEMORAL MODEL

By

Peini Chao

August 2012

Chair: Antonio Pozzi Major: Veterinary Medical Sciences

To compare the bending stiffness, gap motion, cyclic fatigue life and load to failure

of plated dog femora stabilized using either a short or long plate working length

technique. Contralateral femora were randomly assigned to one of two stabilization

techniques: femora plated with 12-hole 2.4 mm LCP using 2 screws per fracture

segment resulting in a long working length or femora plated with the same LCP using 5

screws per fracture segment resulting in a short working length. Constructs stiffness and

yield load in load to failure test were compared between techniques using a paired-t test.

Bending stiffness and gap motion during cyclic testing were compared between

techniques using a 2 x 9 or 2 x 11 (technique x cycle) within subjects repeated measures

ANOVA; a p<0.05 was considered significant.

Construct stiffness did not differ significantly between techniques. Implant failure did

not occur in any of the plated femora during cycling. Mean ± SD of gap motion of long

(0.2 ± 1 mm) and short plate working length constructs (0.1± 0.6 mm) were not

significantly different during cyclic test. Mean ± SD yield load in short plate working

length (654.5 ± 149.0 N) technique were significantly higher than in long plate working

length (452.9 ± 63.4 N) technique. Plate working length had no effect on stiffness, gap

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motion and resistance to fatigue in our testing model. Although short plate working length

tends to increase the resistance to failure; the yield loads for both stabilization

techniques were within physiologic range.

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CHAPTER 1 INTRODUCTION

Over the past two decades, there has been a paradigm shift regarding the approach

to internal fixation of long bone fractures with bone plates. The prerequisite for open

anatomic reduction and rigid stabilization has given way to less invasive application of

more flexible constructs with bridging plates 1-3. The use of locking technology allows the

plate to function as an internal fixator 4-9.Surgeons can choose among a variety of

implants to allow a more or less flexible bone-plate construct 8,10-29. Furthermore, the

design and type of plate can play a primary role in fracture reduction 12,26,29-32. Whereas

plates with a compression or neutralization function require precise reconstruction and

rigid stabilization, plates applied in a bridging fashion circumvent the need for anatomical

reduction33. Bridging plates are often applied utilizing indirect reduction techniques,

which mitigate the degree of iatrogenic trauma while preserving fracture vascularity

1,2,4,32,34-37. Understanding the basic biomechanical principles of surgical stabilization of

fractures is essential for developing an appropriate preoperative plan as well as making

prudent intra-operative decisions.

The objective of this chapter is to provide basic biomechanical knowledge essential

to understand the complex interaction between the mechanics and biology of fracture

healing. It is clearly understood that careful soft-tissue handling is very important in

preserving blood supply to the injured bone. However, the type of healing and the

outcome can be influenced by several mechanical factors, which depend on the

interaction between bone and implant 31,38-41. The main objective for utilizing less

invasive fracture stabilization techniques is to optimize the healing potential by achieving

a symbiotic balance between the biologic and the mechanical factors of fracture fixation.

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Thus the surgeon should understand the mechanical principles of fracture fixation and

be able to choose the best type of fixation for each specific fracture.

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CHAPTER 2 BACKGROUND

Basic Mechanics of Materials

Force, Deformation, Stress and Strain

The strength of a material depends on its ability to resist failure from an applied stress.

Stress is the force acting on an area, and can be compressive, tensile or shear stress.

The unit for stress is Newtons per square millimeter (N/mm2) 42. When stress is applied

to an object, deformation may occur. Thus, the term deformation is used to describe the

change in shape or size of an object due to an applied force 42. Depending on the size,

shape, material of the object, and the force applied, various type of deformation may

occur. Elastic deformation is reversible: an object may deform when subjected to an

applied load, but the object returns to its original shape once the load is released. Plastic

deformation in contrast is irreversible and the object does not return to its original shape

once the applied load is released. Another type of deformation, unique to ductile metals,

is metal fatigue. This phenomenon describes the progressive formation of cracks which

develop in a material subjected to numerous cycles of elastic deformation 43. The

different phases of deformation of an object are readily illustrated by plotting a

load-deformation curve (Figure 2-1), which shows the relationship between stress (force

applied) and strain (deformation) of an object 42. The elastic range of the curve ends

when the object reaches its yield strength and begins to undergo permanent plastic

deformation. If continued load is applied, material failure may occur in the form of a

fracture, or the object may just continue to undergo further plastic deformation. When

applying these concepts to fracture fixation, a bone plate should function within the

elastic region, and should not be subjected to loads which exceed the plate’s yield

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strength. Therefore, yield strength is a very useful parameter for comparing the

mechanical properties of different plates; however, this information is most meaningful

when the applied load in vivo is known.

The term strain is used in order to give a more standardized and quantified description

of material deformation resulting from an applied stress. Strain is defined as the ratio

between the measured changes in length during loading to the original length. Strain

refers to a change in shape of a specified segment which undergoes either elongation or

shortening. Strain is a unitless ratio (length over length), but is commonly reported in

units of microstrain, so that a strain of 0.01 (1%) would be 10,000 microstrain.

Interfragmentary strain is a term used to describe the mechanical enviroment within a

fracture gap subjected to axial loading 44,45. Interfragmentary strain is defined as the

relative change in the fracture gap divided by the original width of the fracture gap

1,41,44,45.

Stiffness

A material’s or a structure’s stiffness defines its ability to resist deformation resulting

from an applied force 42. For so-called linear elastic materials (such as bone), the elastic

region of the load-displacement curve is linear, because the deformation is directly

proportional to the applied load (Figure 2-2). The slope of the linear portion of the

load-displacement curve is known as Young’s modulus, or the modulus of elasticity and

is a measure of the stiffness of the material. Elasticity is a characteristic of a material or

object to return to its previous shape after an applied load is released. Plasticity in

contrast to elasticity describes residual deformation of a material or object as a result of

loading and is an unrecoverable status. More elastic materials can usually sustain

considerable plastic deformation, while more brittle materials will fracture at relatively low

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loads. As an example, the load-deformation curves of the same object with 3 different

material such as bone, glass and metal would look significantly different from each other

(Figure 2-2) The more brittle rigid body such as bone undergo minimal plastic

deformation before reaching the failure point, while metal has a linear elastic portion of

the curve, indicating linearly elastic behavior. The long plastic region of the metal

indicates that this material deforms extensively before failure.

Another concept that helps elucidate the biomechanical characteristics of orthopedic

implants such as plates, screws, intramedullary pins and interlocking nails is the area

moment of inertia 46. The bending stiffness of an object (such as an orthopedic implant)

depends on the bending moment (the force applied to the plate), which is the product of

the elastic modulus of the object material and the area moment of inertia of the

cross-section of the object (Figure 2-3). The area moment of inertia describes the

capacity of the cross-sectional profile of an object to resist bending. The greater the area

moment of inertia, the less a structure will deflect (higher bending stiffness) when

subjected to a bending load. Area moment of inertia is dependent on an object’s

cross-sectional geometry (Figure 2-3). The further the object’s mass is located from the

neutral axis, the larger the moment of inertia. For this reason area moment of inertia is

always considered with respect to a reference axis, in the X or Y direction, which is

usually located at the center of an object’s cross section. The area moment of inertia of

an object having a rectangular cross-sectional profile, such as a plate, can be derived by

the equation bh3/12 where b is the base and h is the height. The base dimension is

oriented parallel to the axis of the moment of inertia and height is defined as the

dimension parallel to the direction of the applied load. Thus the location a plate is

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positioned on a bone and the plate’s orientation to applied bending loads can have a

profound effect on a construct’s bending stiffness (Figure 2-3). This effect becomes even

more important when fractures are not anatomically reconstructed and plates are applied

in bridging fashion. Understanding the area of moment of inertia is important when

comparing the mechanical properties of implants of different shape and size but similar

metal, such as in case of an interlocking nail, a bone plate and a plate-rod construct

(Figure 2-4). As shown in the calculation, the interlocking nail has the largest area of

moment of inertia because of its large diameter. It should also be noted that the area of

moment of inertia of an intramedullary pin occupying 40% of the medullary canal has a

significant contribution to the total area of moment of inertia of a plate-rod construct

(Figure 2-4), justifying its recommendation as bridging implants for comminuted fractures

18,47. Another approach to fixation of a fracture with a gap is to increase the size of the

plate. As shown in the calculation (Figure 2-4), a 3.5 mm broad LCP has an area of

moment of inertia 3 times larger than a 3.5 mm LCP and almost twice larger than a 3.5

mm LCP combined with an intramedullary rod.

Although the stiffness of a plate is an important predictor of the implant’s behavior

under applied load, the mechanical properties of the combined plate-bone construct are

more relevant to predict the type of fracture healing 14. For this reason, it is important to

distinguish between implant stiffness, structural stiffness of the construct and stiffness

across the fracture gap 8,48-50. The construct stiffness is determined by numerous

variables, including the plate’s composition and geometry, the distance between plate

and bone surface, plate length, type of screws and the plate working length

4,8,10,16,27,28,51-58. The plate working length is defined as the distance between the proximal

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and distal screws closest to the fracture 34. The gap stiffness is derived from the

load-displacement curve describing the mechanical behavior of the tissue forming at the

fracture gap. Interfragmentary strain is defined as the relative displacement of the

fracture gap ends divided by the initial fracture gap width 44,59. For this reason the initial

fracture gap is an important factor determining the stiffness of the construct. The

relationship between gap strain and fracture healing has been extensively studied and

will be discussed in the next section.

Fatigue Failure

Mechanical failure of plates can be broadly divided into three categories: plastic, brittle

and fatigue failure. Plastic failure describes failure of an implant to maintain its original

shape, resulting in a loss of reduction and alignment and clinical failure. Brittle failure, an

unusual course of implant failure, results from a defect in design or metallurgy. Fatigue

failure occurs as a result of repetitive loading at an intensity considerably below the

normal yield strength of the implant 42,43. Cyclic loading can lead to the formation of

microscopic cracks that can propagate until these cracks reach a critical size which

cause sudden failure of the implant. Although the propagation of the microcracks can

take a considerable amount of time, there is typically very little, if any, warning preceding

ultimate failure. Crack formation is commonly initiated at a “stress concentrator” or a

“stress riser” such as a scratch on the plate or at a location where there is a change in

the plate’s cross sectional geometry, such as a screw hole. The stress that is focused in

these areas can be relatively higher than the average stress of the whole construct.

Therefore, local material failure can occur at one of these stress concentrators and

eventually propagate through the implant 43. The number of cycles required to cause

fatigue failure decreases as the magnititude of the stress increases. Fatigue failure is a

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genuine concern following fracture stabilization because of the high number of repetitive

loads that implants are subject to during the post-operative convalescent period.

Therefore, every surgeon must be cognizant that they have entered the proverbial race

between fatigue failure of the implant and healing of the fracture, when repairing any

fracture.

Cyclic testing is useful to detect the performance of an implant in resisting fatigue

failure. In general, a predetermined stress or strain limit will gradually be reached while

applying a cyclic bending moment or tensile/compressive force on a construct. The

principle of this type of testing is to determine the total number of loading cycles that a

particular construct can withstand before failing. A construct’s fatigue behavior can be

described in an S/N curve; in which the stress to failure, S, is plotted against the number

of cycles to failure, N (Figure 2-5) 42. A construct’s failure point is termed “allowable

stress”. Typically, a construct subjected to a small applied-stress can withstand a large

numbers of cycles and vice versa. However, the number of cycles to failure at a constant

stress level can be affected by many factors such as the material and the geometry of

the construct or the surrounding environment of the implant and the bone.

Applied Biomechanics

Biomechanics of Fracture Healing

Numerous studies have shown that the mechanical conditions affecting the fracture

site, principally the stability afforded by the fixation and the width of the fracture gap,

influence callus formation during the healing process 28,44,60-70. The process of bone

healing is dependent on numerous interactions between biologic and mechanical factors.

The type of injury, the location and configuration of the fracture, the magnitude of load

acting on the fracture as well as systemic factors, all play a role in the type and efficiency

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of bone healing 41,60,71. Two principle concerns are whether there is adequate blood

supply as well as effective stability to allow a fracture to obtain union. If the local

circulation is adequate to support fracture healing, the pattern of bone healing is then

dependent on the surrounding biomechanical environment 41,44,60,65,68-71. Several

mechano-regulation theories of skeletal tissue differentiation have been developed that

predict many aspects of bone healing under various mechanical conditions 59,66-70,72. The

theory proposed by Perren is based on the interfragmentary strain present in the fracture

gap 44,59. This theory suggests that the type of tissue formed in a healing fracture gap is

dependent on the strain environment within the gap. The tissues that are stressed

beyond their ultimate strain could not form in the gap. If interfragmentary strain exceeds

100%, non-union will occur. Gap strains between 10 and 100%, allow for fibrous tissue

formation. Strains between 2 and 10% allow for cartilage formation and subsequent

endochondral ossification. Strains of less than 2% allow for direct bone formation and

primary fracture healing. Perren proposed that once any tissue formed, it would

progressively stiffen the fracture gap. In turn, the tissue formed in the gap would lead to

lower strains, which would allow formation of the next stiffer tissue and the cycle would

repeat until bone formed within the gap. An alternate theory relating mechanical stimulus

to fracture healing was proposed by Carter and Blenman which predicts that tissue

differentiation within the fracture gap is dependent on the magnitude and the type of local

stress, including hydrostatic pressure and octahedral shear stress 68,70,73. This theory

purports that the vascular supply to the tissues at the fracture site is the primary factor

determining tissue differentiation. With a good circulation in tissues, Carter and Blenman

proposed that fibrocartilage will form if high hydrostatic compressive stresses are

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present. In an analysis of fracture healing, Carter and Blenman correlated compressive

hydrostatic stress with cartilage formation (chondrogenesis) whereas low hydrostatic

stress corresponded to bone formation (osteogenesis) 68-70. However, the relationship

between the ossification pattern and the loading history was described only qualitatively

and not quantitatively. More recently, Claes and Heigele have proposed and tested the

quantitative tissue differentiation theory, which relates the local tissue formation to the

local stress and strain in a fracture gap 66. The results regarding the global strain and

hydrostatic pressure fields correlate with the principal results of Carter and Blenman. In

contrast to Carter and Blenman’s work 68,70,73 the quantitative tissue differentiation theory

is based on the assumption that new bone formation only occurs on existing bony

surfaces and under defined ranges of strain and hydrostatic pressure. The tissue

differentiation hypothesis predicts intramembranous bone formation will proceed once

interfragmentary strain become less than 5% while endochondral ossification can occur

at interfragmentary strains approximating 15% in a diaphyseal fracture, which obtain

union by secondary bone healing 45,66. Another recent theory on mechanobiology of

fracture healing proposed a model dependent on two biophysical stimuli: tissue shear

strain and interstitial fluid flow 67. The rationale for this approach is that fluid flow

increases the biomechanical stress and deformation on the cells above that generated

by the strain of the collagenous material 67.

Although Perren’s theory on interfragmentary strain is important to understand the

concept of tissue mechanobiology at the fracture gap, several studies have

demonstrated that gap strain higher than 2% is tolerated and that the strain patterns

within a fracture gap are heterogenous 74,75. It is well accepted that interfragmentary

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movement is the most important biomechanical factor in fracture healing, but it is not

known which is the optimal range for callus formation and bone healing.

Bone Healing under Conditions of Absolute and Relative Stability

The term stability is defined as the load-dependent displacement of the fracture

surfaces. Stability in osteosynthesis covers a spectrum from minimal to absolute.

Absolute stability is only present when there is no displacement of the stabilized fracture

segments under loading (Figure 6). Absolute stability is achieved by 1) applying a

compressive pre-load that exceeds the traction force acting at the segments, and by 2)

counteracting the shear forces acting on the fracture surfaces with friction. The

elimination of relative motion between the bone segments results from the application of

interfragmentary compression and requires anatomic reduction 9. Placement of a lag

screw is an excellent example of a fixation that can provide absolute stability (Figure

2-6). In vivo experiments have shown that a lag screw can produce high compressive

forces (>2,500 N) across a fracture 76. While absolute stability was originally thought to

be necessary for successful management of most fractures, current thinking suggests

that absolute stability is only obligatory when stabilizing articular fractures and only when

interfragmentary compression can be achieved without inducing excessive iatrogenic

damage to blood supply and surrounding soft tissues 9. Limiting soft tissue trauma is an

essential tenant of any fracture repair even when performing a direct open reduction,

efforts should be made to minimize iatrogenic trauma to the regional soft tissues and the

periosteum.

Fractures stabilized under conditions of absolute stability will heal by primary or direct

fracture healing, if anatomically reduced 77-79. Since there is no motion at the fracture

site, there will be negligible callus formation. The fracture heals through the formation of

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osteonal cutting cones and Haversion remodeling of the compressed cortical bone 78,79.

Direct bone healing can be further subdivided into two types based on the width of the

fracture gap. Contact healing occurs when the ends of the bone segments are in direct

contact, the gap between the two bone segments is less than 0.01 mm and when

interfragmentary strain is less than 2% 78. If the fracture gap is larger but does not

exceed 1 mm, and an interfragmentary strain again is less than 2%, gap healing will

occur, in which intramembraneous bone will be formed directly in the fracture gap 44. In

both types, a process called Haversian remodeling begins with osteoclastic resorption

that result in resorption cavities formed by groups of osteoclasts also called a cutting

cone 79. Bone resorption is followed by osteoblast activity. The osteoblasts line the

resorption cavities and produce layers of new bone. The resorption cavity is filled in with

new bone to form a new osteone. Gap healing results from the development of lamellar

and cortical bone forming from granulation tissue in small gaps 78,79. Intramembranous

bone formation occurs during direct bone healing; the surrounding environment can

impose up to 5% strain as long as it allows the differentiation of mesenchymal cells into

osteoblasts..

Relative stability is a condition where an acceptable amount of interfragmentary

displacement compatible with fracture healing is present (Figure 2-7) 80. Relative stability

involves placement of implants that provide somewhat flexible fixation which allow an

acceptable degree of fracture segments displacement. Fixation modalities that can be

employed to provide relative stability include plates, interlocking nails and external

fixators applied in bridging fashion to span a bone defect 71,81.

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Relative stability provides a mechanical environment which promotes indirect or

secondary bone healing 65,80. Indirect bone healing is very similar to the embryological

bone development and occurs via both endochondral and intramembranous ossification

65,80,82. The healing process by formation of callus can be divided in 4 stages:

inflammation, soft callus, hard callus and remodeling 82,83. Mineralized cartilaginous

callus develops at the ends of the fracture segments (gap callus), along the medullary

canal (medullary callus) and on the outer cortex (periosteal callus) 82,83. The majority of

the vascular circulation to the callus is derived from the surrounding soft tissues 84.

Therefore, surgical techniques that preserve the soft tissue envelope adjacent to the

fracture are advantageous and promote fracture healing.

The indications for using techniques that achieve absolute or relative stability differ

according to fracture location, fracture configuration, soft tissue conditions and

vascularity of the bone. Simple transverse, spiral or oblique fractures that can be readily

anatomically reconstructed, are good candidates for anatomic reconstruction and

compression or neutralization plating. More complex comminuted fractures should be

treated with bridging fixation. Articular fractures should be anatomically reduced and

stabilized with fixation that generates interfragmentary compression, such as lag screws

71. It is always important to consider whether it is possible to implement anatomical

reconstruction when choosing the type of fixation. For example, fractures that may

initially appear as simple, reconstructable fractures may instead have fragments that are

too small for anatomic reconstruction. In these cases an open or closed indirect

reduction technique, and a bridging stabilization technique may be indicated. Because

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the success of the technique depends on the precision of the reduction, critical

preoperative planning should always be performed.

Factors Affecting Stiffness of the Plate-Bone Construct

The stiffness of the bone-plate construct is a major determinant of the mechanism and

progression of bone healing 1,28,50,68,69. There are several parameters in addition to the

material properties of the implants that need to be considered when applying a bone

plate. Understanding the effect of plate type, size, length, position, screw type and screw

placement is important because successful fracture healing depends on appropriate

fixation stability 53,55,57,58,85-88. Furthermore a multitude of plate types and concepts have

been described and proposed in the last decade, in an attempt to decrease the

complications and improve the reliability of bone plating. The development of new

implants and techniques have followed a shift in emphasis of AO philosophy from

obtaining anatomic reconstruction and absolute stability to obtaining anatomic alignment

and appropriate stability utilizing more atraumatic application techniques 9,71. Concurrent

with this change in emphasis in internal fixation, newer implant systems such as internal

fixators, locking plates or angular stable devices, have been developed to improve

bone-plating technique 9. Understanding the mechanical properties of locking plates and

conventional plates is important for choosing an appropriate implant system.

Choosing the type of plate: locking versus non-locking plates

Gautier and Sommer recently presented prudent guidelines that may improve the

individual learning curve of surgeons who are less familiar with locking plates 31.

However, it is important to understand the concepts behind these recommendations for

successful utilization of the vast choice of plates available 10,27. There are distinct

principal biomechanical differences between bridging plates and locked internal fixators

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with regard to load transfer through a fractured bone. In conventional compression plate

constructs or nonlocking bridging plate constructs, fixation stability is limited by the

frictional force generated between the plate and the bone. This force is created by axial

screw forces and the coefficient of friction between the plate and the bone 8,89. If the

force exerted on the bone while the patient is ambulating exceeds the frictional limit,

relative shear displacement will occur between the plate and the bone, causing a loss of

reduction between the bone fragments, or loosening of the screws, or both. Conventional

plates, including dynamic compression plates (DCP) 90 and limited contact dynamic

compression plates (LC-DCP) 91 allow compression of bone fragments utilizing the

dynamic compression holes. In a transverse fracture that has been anatomically

reduced, stability can be further increased by utilizing the plate to generate

interfragmentary compression between the ends of the fracture segments when the

screws are inserted eccentrically at the end of the oval hole (away from the fracture), the

lower hemi-spherical part of the screw head will contact the dynamic compression incline

of the compression hole. This interaction between the screw head and the compression

incline results in sliding of the screw down the incline hole in the plate, producing

compression of the ends of the fracture segments during screw tightening 90,91.

Locking plates differ from non-locking plates because stability is not dependent on the

frictional forces generated at the bone-plate interface. The first plate that functioned as

an internal fixator (Zespol system) was developed in 1970 in Poland 92. Since then, a

number of locking plates have been developed that utilize the concept of angular

stability. These implants consist of a plate and locking head screws which together act

as an internal fixator. Locking the head screw into the hole confers axial and angular

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stability of the screw, relative to the plate. Because the stability of the construct does not

depend on frictional forces generated between plate and bone, the bone-screw threads

are unlikely to strip during insertion. The fixed-angle connection between the screw and

the plate clearly affords improved long-term stability. Plate failure by “pull-out” is unlikely

because the screws cannot be sequentially loaded or pulled out 9,25,93.

Locking plates have both mechanical and biological advantages. The periosteal blood

supply beneath the plate is not compromised because compression between plate and

bone does not occur. Preservation of the periosteal vasculature may improve healing

and decrease the risk of cortical bone necrosis and infection 81. Another advantage is

that the plate does not need to be perfectly contoured, because the bone is not “pulled

towards” the plate while tightening the screw. For this reason, locking plates are often

used for minimally invasive plate osteosynthesis (MIPO), which involves closed

reduction and percutaneous fixation of the fracture 34 In press,35,94,95. Several different

locking plate systems are available. Some plates may have combination holes that allow

placement of a locking screw or a conventional non-locking screw in either a

compressive or neutralize fashion or a locking screw 9,96,97.

Several biomechanical studies have compared locking and non-locking plates in

dogs. These studies have conflicting results. While some studies demonstrated that

locking plate constructs were stiffer than non-locking plate constructs when tested in

axial compression, torsion, and bending 16,22,51,89,96,98-104, others did not find any

significant differences between the two 12,20,26,29,54,86,98,105-108. The most consistent finding

has been that locking plates perform better than non-locking plates in osteoporotic bone

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4,89,99,109. The biomechanical advantages of locking plates may be less evident in normal

bone.

Choosing the length of the plate

The selection of an appropriate length plate is a very important step in the

preoperative plan. Appropriate plate length is dependent on the location and

configuration of the fracture as well as the intended functional application of the plate. In

bridge plating, longer plates lower the pullout force acting in screws due to an

improvement of the working leverage for the screws and better distribution of the

bending forces along the plate 31. The theoretical advantage of using a longer plate

without placing screws in the center portion of the plate is supported by several

biomechanical studies. Sander et al compared three different plate lengths, six-, eight- or

ten-hole 3.5 mm dynamic compression plates fixed on dog cadaveric ulnae, tested in

four-point bending to failure. The results revealed that ten-hole plates with four screws

(widely spread on the fracture fragment), failed at higher peak loads than six-hole plates

with six screws, supporting the recommendation tha longer plates with fewer screws

provide superior bending strength than shorter plates with a greater number of screws 88.

In another study, Weiss evaluated eight- and ten-hole 3.5 mm locking compression

plates used to stabilized human cadaveric ulnas. This study found that 10-hole plates

secured with two non-locking screws placed in a near-far configuration on either side of

the fracture demonstrated an increased yield strength compared to 8-hole plates with the

same number of screws and configuration in four-point bending to failure 86. Iatrogenic

trauma associates with the open application of a long plate can be substantially mitigated

by using less invasive application techniques such as MIPO.

27

Two values have been used to determine the length of the plate to be used. The plate

span ratio is a quotient derived by dividing plate length by the segmental length of the

fracture. Based on guidelines developed for fracture fixation in man, the plate span

should be more than 2 to 3 in comminuted fractures and more than 8 to10 in simple

fractures 9. Plate-screw density is the quotient derived by dividing the number of screws

inserted by the number of holes in the plate. Empirically, values below 0.4-0.3 when

applied in simple fracture and value below 0.5-0.4 when applied in comminuted fracture

have been recommended 9,31. These guidelines were formulated for the application of

plates in human patients, and need to be evaluated in dogs.

Effect of the position of screw placement in the plate

In comminuted fractures which have not been reconstructed, stress is distributed over

the fracture gap and is dependent on the number and location of screws placed 9,53 in

addition to other factors. The lowest stress in the plate occurs when the screws are as

close as practical to the fracture 53. This, however, leads to the highest axial stiffness as

well as very small interfragmentary movements and strains beneath the plate. It has

been recommended to increase the plate working length to reduce axial stiffness of a

plate-bone construct 9,31,110; however, previous mechanical studies have yielded

conflicting results 53,87,111. Based on mechanical tests performed in our lab 34 we suspect

that the variability in the results amongst reported studies might be attributed on how the

plate is applied to the bone. In constructs which utilize non-locking plates, the contact

between the plate and the bone segments causes the bending moment to concentrate

between the end of the bone segments regardless the positioning of the screws.

Therefore, the functional plate working length does not correspond to the distance

between the screws placed closest to the fracture gap, but rather to the length of the

28

fracture gap. In contrast, the physical offset of a locking plate which is applied without the

bone and plate in intimate contact enables a locking plate to bend along the entire

segment of the plate between the two most centrally positioned screws.

More recent strategies to decrease the stiffness of locking plate-bone constructs

include new designs of locking screw that allow to increase the fracture gap micromotion

with axial loading 112-114. The goal of this novel approach to locking plate is to promote

more reliable healing, and prevent late failures observed in several clinical studies in

people 115-117. The far cortex locking screw has a smooth shaft with threads at the tip that

achieve purchase in only the far cortex 50,112,113,118. The smooth shaft of this screw

decreases the stiffness of the plating construct and allows greater callus compared to

standard locked implants 112. Another screw design attempting to combine the

advantages of locking screws and controlled axial micromotion is the dynamic locking

screw 114. This dynamic locking screw is composed of an outer sleeve with threads that

engage the bone and an inner pin with threads that lock to the plate. By allowing motion

between inner pin and the outer sleeve, the dynamic compression screws reduced the

axial stiffness by 16% 114.

29

Figure 2-1. A load-deformation curve shows the relationship between stress (force applied) and strain (deformation) of an object. The elastic range (A-B) of the curve ends when the object reaches its yield strength (B) and begins to undergo permanent plastic deformation (B-C). If continued load is applied, material failure may occur in the form of a fracture, or the object may just continue to undergo further plastic deformation.

30

Figure 2-2. The load-deformation curves of the same object with 3 different material: bone, glass and metal. More elastic materials can usually sustain considerable plastic deformation, while more brittle materials will fracture at relatively low loads. The more brittle rigid body such as bone undergo minimal plastic deformation before reaching the failure point, while metal has a linear elastic portion of the curve, indicating linearly elastic behavior. The long plastic region of the metal indicates that this material deforms extensively before failure.

31

Figure 2-3. The area moment of inertia describes the capacity of the cross-sectional profile of an object to resist bending. The greater the area moment of inertia, the less a structure will deflect (higher bending stiffness) when subjected to a bending load. Area moment of inertia is dependent on an object’s cross-sectional geometry.

32

Figure 2-4. Area moment of inertia is important when comparing the mechanical properties of implants of different shape and size but similar metal, such as in case of an interlocking nail, a bone plate and a plate-rod construct ( As shown in the calculation, the interlocking nail has the largest area of moment of inertia because of its large diameter. It should also be noted that the area of moment of inertia of an intramedullary pin occupying 40% of the medullary canal has a significant contribution to the total area of moment of inertia of a plate-rod construct, justifying its recommendation as bridging implants for comminuted fractures. Another approach to fixation of a fracture with a gap is to increase the size of the plate. As shown in the calculation, a 3.5 mm broad LCP has an area of moment of inertia 3 times larger than a 3.5 mm LCP and almost twice larger than a 3.5 mm LCP combined with an intramedullary rod.

33

Figure 2-5. A construct’s fatigue behavior can be described in an S/N curve; in which the stress to failure, S, is plotted against the number of cycles to failure, N. Typically, a construct subjected to a small applied-stress can withstand a large numbers of cycles and vice versa.

34

Figure 2-6. Absolute stability is only present when there is no displacement of the

stabilized fracture segments under loading. The elimination of relative motion between the bone segments results from the application of interfragmentary compression and requires anatomic reduction.

Figure 2-7. Relative stability is a condition where an acceptable amount of

interfragmentary displacement compatible with fracture healing is present.

35

CHAPTER 3 EXPERIMENT

Effect of Plate Working Length on Plate Stiffness and Cyclic Fatigue Life in a Cadaveric Femoral Gap Model Stabilized with a 12-hole 2.4 mm LCP

There are several factors that can affect the fatigue life of plated fracture construct.

In addition to the mechanical properties of the isolated plate, the location, type and

complexity of the fracture influence the load acting on the plate 8,17,31,55,85,88. Bridging

plates are often utilized in comminuted fractures to span a large segment of the fractured

bone which is not reconstructed, having a long segment of the plate unsupported by

screws 110. The distance between the proximal and distal screw in closest proximity to

the fracture is defined as the “working length” of the plate. Plate working length has been

shown to influence construct stiffness, plate strain and cyclic fatigue properties of the

plated constructs 57,85,87,111,119. The lack of load sharing between the stabilized bone

segments and the implants increases the risk of cyclic fatigue and potentiates early

failure of the implant 53,87. A mechanical study evaluating the mechanical endurance of

human femora stabilized with 14-hole broad 4.5 mm locking compression plates (LCP)

found that constructs with load sharing resisted 20 times more cycles than the constructs

with an 8 mm segmental diaphyseal gap 120.

Controversy still exists regarding the effect of plate working length on stiffness and

resistance to fatigue failure 53,58,85. Stoffel et al. reported that increasing the plate working

length by omitting one screw placed adjacent to the fracture in each of the major fracture

segments made a locking plating construct nearly twice as flexible when loaded in axial

compression and torsion 53. In contrast, Field et al. reported that omitting two screws

proximal and distal the fracture had no significant effect on either bending or torsional

stiffness of a conventional plate construct in a comparable bridge plating configuration 58.

36

Studies evaluating the effect of working length on the cyclic fatigue properties of plated

constructs have yielded inconsistent results 53,85,87. Stoffel et al. found that the constructs

with a shorter plate working length were more resistant to cyclic loading 53. Maxwell et al.

reported similar results when screw placement was evaluated in 3.5 mm dynamic

compression plates and limited contact dynamic compression plates applied in a fracture

gap model 111. Recent studies comparing the cyclic fatigue of plates applied with a short

and long working length found that the constructs with a longer working length withstood

more cycles before failure, although the differences between stabilization techniques

were not significant 85,87.

The purpose of the present study was to evaluate the effect of plate working length

on construct stiffness, gap motion and resistance to cyclic fatigue in dog femora with a

simulated fracture gap stabilized using a 12-hole 2.4 mm LCP. Our first hypothesis was

that constructs with a longer plate working length would be more flexible and have

greater gap motion than constructs with a shorter plate working length. The second

hypothesis was that constructs with a longer plate working length would withstand more

cycles before succumbing to fatigue failure than constructs with a shorter plate working

length. Materials and Methods

Specimen Preparation

This study was approved by the University of Florida Institutional Animal Care and

Use Committee. Ten pairs of femora were collected from adult dogs (mean ± SD: 17.9 ±

1.9 kg) that were euthanized for reasons unrelated to the study. Once all the soft tissues

were removed from the femurs, the bones were radiographed in order to exclude

specimens with osseous abnormality. The femora were then wrapped in saline-soaked

37

towels and stored at -20℃ until testing, at which time the bones were thawed to room

temperature. One femur from each dog was randomly assigned to the short plate

working length stabilization group. The contralateral femur was assigned to the long

plate working length stabilization group.

A 12-hole 2.4 mm LCP1 was contoured and clamped temporarily to the lateral surface

of each femur. Two lines, 10 mm apart were marked on the mid-diaphyseal region of the

bone adjacent to the central 2 holes of the plate. The plate was removed, and an

oscillatory saw was used to initiate two mid-diaphyseal partial osteotomies in the lateral

cortex of the femur at the marked locations. In the femora assigned to the long plate

working length stabilization group, the plate was affixed with one 2.4 mm locking and one

2.4 mm cortical screw placed at the ends of the plates, leaving eight empty holes in the

middle of the plate (Figure 3-1). The femora assigned to the short plate working length

stabilization group were fixed with one 2.4 mm locking and four 2.4 mm cortical screws in

each femoral segment, leaving two empty holes in center of the plate (Figure 3-1). All

screws had bicortical purchase and were placed in the same order. Cortical screws at

each end of the plate were placed first, followed by the locking screw in the second hole

and the remaining cortical screws. All screws were placed and hand-tightened by an

experienced board-certified surgeon (AP) in accordance with standard AO-ASIF

techniques 121. Prior to testing, the torque of each locking and cortical screw was

tightened to 0.8 Nm or 0.44 Nm, respectively 122. After applying all screws, osteotomies

were completed from the medial cortex of the femur resulting in a mid-diaphyseal

segmental ostectomy.

1 Synthes USA, Paoli, PA

38

The distal ends of femora were potted using epoxy resin2. A goniometer was used to

ensure that the specimens were potted perpendicular to the testing mold. Two points

located on the cranial and caudal surface of each femur segment adjacent to the fracture

gap were first marked then scribed using a 1.5 mm drill bit, so these points could be

accurately identified with the tip of the articulating arm of a three-dimensional positioning

device3.

Mechanical Testing

All testing was performed using a servohydraulic MTS4. During testing, each

specimen was positioned vertically with the epoxy resin block secured to the machine

using a specially designed fixture. A polyoxymethylene cup simulating the shape of the

acetabulum was secured to the mechanical test system and used to load the femora. In

order to measure gap displacement of the bone segments, three points on each bone

fragment were digitized using the Microscribe while the limb was preloaded to 10 N.

A cyclic axial compressive load from 0 N to 100% body weight (mean ± SD: 175.5 ±

18.6 N) was applied at 2 Hz, while displacement data was recorded at 100 Hz throughout

testing. Testing was paused at 1000, 2000, 5000, 10000, 20000, 50000, 100000,

120000, and 18000 cycles to obtain gap displacement measurements as each specimen

was loaded. If fatigue failure did not occur after 180,000 cycles, the specimen was

loaded in axial compression at a rate of 1 mm/seconds until failure was achieved.

Displacement data was recorded at 100 Hz while specimens were loaded to failure.

2 Bondo Corp, Atlanta, GA

3 Microscribe 3DX digitizing arm (Immersion Corp., San Jose, CA)

4 MTS 858 Mini Bionix II, MTS Systems Corp, Eden Prairie MN

39

Failure was defined as breakage or plastic deformation of the implants or any visually

detected loosening of screws in the plate or bone. Video was recorded during testing to

identify the mode of failure for each construct.

Data and Statistical Analysis

Displacement values were recorded at 1000, 2000, 5000, 10000, 20000, 50000,

100000, 120000, and 18000 cycles. A load-displacement curve was calculated by

plotting the load versus displacement data for each specimen for each time point during

cyclic testing in a scatter graph using spreadsheet software5. A best fit trend line, straight

line slope, and corresponding R2 value were determined for each load-displacement

curve; using a sum of least squares method for the linear elastic region of the

load-displacement curve using the same software. The slope of this line defined the

stiffness value and was reported in units of N/mm.

Relative gap displacement of two bone segments in the sagittal and frontal plane was

analyzed using classical principles of rigid body mechanics 123. Gap displacement of

each construct was also calculated to compare whether there was any significant

difference between short and long plate working length stabilization techniques.

Calculations were performed using a custom-written computer program6.

A separate load-deformation curve was determined similarly from the initial linear

segment of load to failure test for each construct. The elastic limit or yield load was

defined as the loading value at which the linear phase of the curve terminated 124. Mean

5 Microsoft Office Excel 2003, Microsoft Corporation, Redmond, Washington

6 MATLAB®, MathWorks Corporation, Natick, Massachusetts

40

values and standard deviations of stiffness, gap displacement and yield load were

determined for both the short and long plate working length stabilization techniques.

Stiffness values, cranial/caudal translation, medial/lateral translation of two bone

segments and gap displacement measured during cyclic testing were compared

between the short and long plate working length stabilization techniques using a 2 x 9

(technique x cycle) within subjects repeated measures ANOVA. When appropriate a post

hoc Bonferroni procedure7 was used to adjust for multiple pairwise comparisons. The

yield load determined during load to failure testing was statistically compared between

the short and long plate working length stabilization techniques using a paired t-test. For

all statistical analyses, significance was set at p < 0.05.

Results

When the constructs were cyclically loaded, there was no significant difference

(p=0.435) in stiffness between the short and long plate working length stabilization

techniques. There was a trend for the stiffness to increase during cycling in both short

and long working length plate stabilization techniques until 50,000 cycles. After reaching

50,000 cycles, the mean stiffness of long working length plate technique decreased

while the mean stiffness of short working length plate technique continued to increase,

although these trends never became significant between stabilization techniques (Figure

3-2). All constructs completed 180,000 cycles of loading without plate or screw

breakage, screw loosening or visible plastic deformation of the implants observed in any

of the constructs.

7 SPSS Statistics 17.0, SPSS Inc., Chicago, IL

41

There were no significant differences in the relative displacement of the bone

segments between stabilization techniques in the sagittal (p=0.448) or frontal plane

(p=0.504). Gap displacement also did not differ significantly within the designated cycles

and between stabilization techniques (p=0.116). The ends of two stabilized bone

segments did not come into contact in any of the constructs during cyclic compression

loading.

During load to failure testing, constructs in the short plate working length

stabilization technique had a significantly (p=0.016) higher yield load (mean ± SD: 654.5

± 149.0 N) than constructs in the long plate working length stabilization technique (mean

± SD: 452.9 ± 63.4 N). The mean ± SD stiffness of the short (223.8 ± 46.7 N/mm) and

long (219.4 ± 80.3) plate working length stabilization techniques during load to failure

testing was not significantly different. Failure patterns of the short and long plate working

length stabilization techniques were distinctively different (Figure 3-3): the proximal

femoral segment bent medially during axial loading in both the short and long plate

working length constructs. The lateral surface of the proximal femoral segment acted as

a fulcrum against the plate and there was no separation of the plate from the lateral

cortex of the femur. The plate elevated from the lateral cortex of the distal femoral

segment in the long plate working length constructs. The plate remained apposed to the

lateral cortex of the femur in the short plate working length constructs.

Review of video obtained during the load to failure testing revealed that eccentric

axial compressive load applied to the femoral head leads to the formation of a bending

moment, with the center of rotation located at the plate bone interface. Bending moment

was defined as the product of applied force (F) amount and the perpendicular distance

42

(d) from the applied force to the axis of rotation (Figure 3-3). The bending moment

increased during the load to failure test as the plate deflected away from the lateral

cortex of the femur. This effect was more pronounced in the long plate working length

constructs.

Discussion

Increasing working length by placing a limited number of screws at each end of the

plate has been recommended as a strategy to decrease construct stiffness and therefore

allow more motion at the fracture gap 9,31,48,53,110; however, previous mechanical studies

have yielded conflicting results 53,57,87,111. Based on our results, we cannot support our

hypotheses that constructs with fewer screws and a longer plate working length would be

more flexible, have greater gap motion and more resistant to cyclic fatigue failure than

constructs with more screws and a short plate working length. We suspect that the

inconsistencies in the results amongst reported studies might be attributed to the manner

in which the plate was applied to the bone. In constructs which utilized non-locking plates

58,111, the contact between the plate and the bone segments causes the bending moment

to concentrate between the ends of the bone segments regardless the positioning of the

screws. Therefore, the functional plate working length does not correspond to the

distance between the screws placed closest to the fracture gap, but rather to the length

of the fracture gap. In contrast, the physical offset of a locking plate which is applied

without the bone and plate in intimate contact enables a locking plate to bend along the

entire segment of the plate between the two most centrally positioned screws 53. If our

suppositions are confirmed, the concept of short and long plate working length should

only be applied to locking plates and should not be applied to conventional non-locking

plates.

43

The results of our cyclic fatigue testing are not consistent with previous studies

which found differences between constructs with short and long plate working lengths

53,111. These contrasting results might be ascribed to differences in the fracture models

employed and methods of loading. In our study the applied load was based on the dogs’

body weight in an attempt to replicate clinical situations. In other studies the upper load

threshold was set to induce implant failure within a predetermined number of cycles 85 or

was established using displacement control based on plate strain 87. Establishing

significance may have been masked in our study because we did not apply high enough

loads as none of our constructs failed during cyclic testing. The magnitude of the load

applied in diaphyseal fracture models is more complex than simply applying a load

equivalent to mean body weight. In human patients which have been instrumented with

distal femoral diaphyseal prosthesis with telemetric strain gauges, load values as high as

three times body weight and bending forces up to ten times body weight have been

recorded 125. Other loading protocols such as progressive loading 112 could have been

used to shorten the test protocol and may have yielded significant differences between

stabilization techniques.

Both the short and long plate working length stabilization constructs sustained

180,000 cycles with the equivalent of full body weight loaded in axial compression

without failure, which approximates 2 to 4 months of weight-bearing during normal

ambulation 126. This finding suggests that both constructs would likely provide acceptable

clinical stability as most fractures are expected to obtain union within 12 weeks of

stabilization 38,81. The yield load for both stabilization techniques ranged from 453 to 655

N, which corresponds to 2 to 4 times the hindlimbs peak vertical force measured from a

44

normal 20 kg dog running at a trot 15,127. Although kinetic values measured with gait

analysis are only an approximation of the forces tolerated by implants during loading, our

results suggest that it is unlikely that either of our constructs would fail catastrophically in

the early postoperative convalescent period.

Cadaveric studies have numerous limitations. When performing cyclic testing

designated to resemble a clinical environment after fracture fixation, the loading plane

should be considered 128. We selected an offset axial loading to simulate loading of a

plated femoral fracture. Our testing methodology had the limitation of being isolated to a

single plane, without considering more complex forces such as a combination of bending

and torsional forces. In the diaphyseal region of the femur, however, axial and bending

forces predominate and these forces were replicated in our testing protocol 125. Another

limitation was that we did not test a construct which exclusively utilized locking screws.

This omission limits our ability to make conclusions regarding the effect of working length

of a locking plate employed in its intended application 53,87. A hybrid construct in which

both locking and non-locking screws were used was selected because these constructs

are commonly used in dogs16. Furthermore, previous studies have shown that placement

of a single locking screw in each of the major fracture segments increases construct

stability subjected to axial and torsion loading 16,96,129,130.

We were unable to establish a significant effect of working length on stiffness and

resistance to cyclic fatigue in a fracture gap model plated with 2.4 mm LCP applied in

direct cortical contact with the femur. Our results question the advantage of long working

length in decreasing stiffness of the bone-plate construct and protecting the portion of

unsupported plate at the fracture gap 6,48,131. Neither of the constructs tested in our study

45

failed during cyclic load. Therefore both short and long working lengths constructs would

appear to be acceptable for clinical applications. Placing five bicortical screws in each of

the major fracture segments may not be necessary. Other studies have reported that

fewer, more widely spaced screws increased the bending strength of screw–plate

fixation more effectively than increasing the number of screws 31,55,131. Guidelines

regarding the optimal number of screws will, however, need to be derived from future

mechanical and clinical studies.

46

Figure 3-1. Constructs tested: long plate working length stabilization technique (left) and short plate working length stabilization technique (right). Each plate had a

2.4 mm locking screw (○) placed in the second hole from each end of the

plate. The other screws were 2.4 mm cortical screws (●).

Figure 3-2. Comparison of mean ± SD values of construct axial stiffness (N/mm) loaded to body weight between long plate working length stabilization

technique (□) and short plate working length stabilization technique(■).

Measurements recorded at 1000, 2000, 5000, 10000, 20000, 50000, 100000, 120000, and 180000 cycles. No significant difference was found between the two stabilization techniques at any of the evaluated cycles.

47

Figure 3-3. Photographs of a long and short plate working length construct before (left) and after (right) load to failure testing. The adjacent free body diagrams represent how bending moments were created. Bending moments (Force (F) x Distance (d)) were calculated based on measured scale distances obtained from photographs and the yield load of each construct.

48

CHAPTER 4 CONCLUSION

Plate working length is a major determinant of construct stiffness 53. A guideline of

whether to keep plate working length short or long remains controversial. The

recommendation of keeping at least two or three empty plate holes at the fracture site

were made in order to allow adequate construct flexibility when bridging a fracture 31,131.

Additionally it was theorized that two or three empty plate holes at the fracture site would

decrease plate strain at the fracture gap, therefore decreasing the risk of fatigue failure

9,131. However, several studies have demonstrated that a screw should be placed as

close to the fracture gap as possible in each major fracture segment, to decrease plate

strain and increase the resistance to fatigue failure53,57. Based on our testing results we

cannot support the hypotheses that constructs with longer plate working length

stabilization technique have lower stiffness, greater gap motion and are more resistant to

fatigue failure. No significant effect of working length were found on construct stiffness,

gap motion and resistance to cyclic fatigue in a fracture gap model plated with 2.4 mm

LCP applied in direct cortical contact with dog femur. Both short and long plate working

length stabilization techniques should be able to sustain through post-surgery coalescent

period in dogs.

Our results are inconsistent with previous studies in which they found that longer

plate working length stabilization technique has less stiffness, higher tensile strain and

earlier fatigue failure than short plate working length stabilization technique 53,57,111.

Inconsistent findings may results from different fracture model and loading methods. In

our study we fixed the plates flush on the femoral cortex which probably limited the

effective working length to the portion of the plate bridging the gap. In other words, the

49

working length may equal the size of fracture gap in a non-locking plate-bone construct.

Stoffel et al. applied plates at 2 mm offset from plate to bone, allowing plate bending over

the entire distance between two inner most screws on the plate. Therefore the plating

working length in a locking plate-bone construct truly corresponds to the distance

between the two innermost screws. Stoffel et al. focused on the contribution of the plate

to stiffness of the construct. This result results are similar to other studies that found no

significant difference of constructs stiffness and fatigue resistance between long and

short plate working length stabilization techniques 85,87. The mode of failure and plate

deformation observed in our study was consistent with Hoffmeier et al.’s study. In long

plate working length stabilization constructs, the free plate sections on both sides of the

gap bulged away from the bone during bending. Hoffmeier et al. found that the larger the

working length, the larger the radius of curvature would be and the evenly the stress

would be distributed. We might found the similar result of the stress distribution on the

plate because we have used similar femoral fracture model without leaving distance

between plate and bone as in Hoffmeier et al.’s study.

We applied LCP in hybrid fixation in which one locking screw was placed next to

other cortical screws in each bone segment. The effect of using locking screws in

combination with cortical non-locking screw in LCP applications is multifold: using

locking screws has been reported to increase axial strength and torsion resistance, as

well as increase the remaining torque in nearby cortical screws than all non-locking

screws fixation 16,129,130. Even though there is currently no study comparing the effect of

different plate working length on hybrid plate fixation, one study suggested that hybrid

plate fixation has similar bending strength and higher torsional strength than an

50

all-locked bridge plating construct in the osteoporotic diaphysis 97. Therefore, it is unlikely

that the hybrid plate fixation we used was the reason led to different results between our

study and other all-locked- screw study.

Cyclic loading is the mechanical test that more closely resembles the conditions of

fracture-healing in the clinical patient. However, it presents several limitations and

challenges. First, the loading plane should simulate the in vivo conditions that the implant

would experience in vivo. In our study we selected axial compression loading which

cause mainly bending force on the femur, because it is one of the most commonly force

that femurs encountered in physiological conditions 29,56. The magnitude of load is

another major factor that should be considered when designing a cycling test. To

simulate in vivo condition of loading, clinically relevant forces should be used. In people,

the physiological basis for testing loads is based on telemetrized implants such as

interlocking nails or total knee replacements, which directly measure the forces acting at

the level of the femoral diaphysis or at the knee joint, respectively. Because we lack of

these data for dogs, we chose testing loads based on dog’s body weight in an attempt to

replicate clinical situations 15.

The number of cycles required to test an implant for fracture fixation is controversial.

The presence of callus is a major factor in construct stability because it may develops as

early as 2 weeks following fracture fixation. Fracture callus protects the implant from

failure by providing load sharing, but cannot simulated in a laboratory setting. Our study

design is different than other studies in which fatigue tests were designed either the

upper load threshold was set to induce implant failure within a predetermined number of

cycles 85 or the number of cycles are as high as 1,000,000 cycles 17,53. Some other

51

studies used protocols as control displacement based on plate strain87 or applied

progressive loading 118, aimed at shortening the duration of cyclic fatigue test while also

assuring if significant differences between stabilization groups. Because all constructs

sustained through 180,000 cyclic weight-bearing loading in our study, the loads and the

number of cycles we chose were not high enough for constructs to reach fatigue failure.

Therefore, the significance may have been masked in our study.

The material of the plate is another factor affecting fatigue endurance of plate-screw

constructs. Hoffmeier et al compared the fatigue life between short, medium and long

working length in both stainless steel and grade-2 titanium 4.5 mm distal femoral plates

which bridged over a 10 mm fracture gap on a composite human femur 85. When plates

of different metals were compared, the stainless steel plates sustained higher cyclic load

than the grade 2-titanium plates. For the steel plates, no significant correlations were

found between fatigue strength and different working length groups. For the titanium

plates the fatigue strength was greater in constructs with long plate working length but

this correlation was not significant. Therefore, the effect of plate working length on

fatigue endurance may be greater in titanium plates than in stainless steel plates.

We did not evaluate other factors that affect plate-screw constructs stability such as

the number of screws inserted in each fragment. Studies which evaluated the effect of

different screw numbers on stiffness of plate-screw constructs reported that fewer, more

widely spaced screws increased the bending strength of screw–plate fixation more

effectively than increasing the number of screws 31,55,131. Other studies compared the

effect of different screw positions on plate strain or fatigue resistance 85,111. Therefore,

the effect of screw numbers should be minimal compared to the position of screws.

52

Future studies are necessary to evaluate if it is necessary to place screws in each screw

holes in the plate, especially considering the vascular trauma caused by drilling screw

holes in cortical bone 132.

Stoffel et al made the statement that as long as more than four screws were placed

each bone segment, the torsional rigidity would not be affected by the position of the

screws. Similar with other study suggested that regardless of the spacing of screws and

the plate length, strength in torsion was dependent on the number of screws securing the

plate 53,55. One study also reported that add a single locking screw to an otherwise

non-locking construct would increase the torsion rigidity 16. Therefore, if we tested our

constructs in torsion, the major variable that causes any difference should be the number

of screws rather than plate working length.

Longer plate working length stabilization technique has been shown to result in

higher plate strain in both experimental and computational study models 53,57,119,133. The

authors concluded that screws should be placed as closed to the fracture as possible in

order to decrease plate strain. However, one study evaluated the strain at and around

the fracture site in a gap model stabilized by 12-hole 3.5 mm LC-DCP and DCP, and

reported that the gap strain does not significantly change after screw removal 111. The

other study used a femoral fracture model and found a reduction of plate strain with

increasing plate working length stabilization technique 87. This finding is supported by a

computational simulation showing that a short plate working length produces stress

concentrations in the plate, which leads to fatigue failure 134. We did not measure plate

strain during test; however, we did not find any significant difference in gap motion during

cyclic testing. Based on our results, it is likely that we would not find any difference in gap

53

strain because gap strain is affected by gap motion. After load to failure test, both short

and long plate working length stabilization techniques have similar bending fulcrum at

the end of the proximal bone segment, which is likely the area where we would find the

greatest plate strain. However, the strain distribution along the entire plate between two

plate working length stabilization techniques during cyclic loading test would need further

testing to be determined.

Both short and long plate working length stabilization constructs should be

acceptable for clinical applications, because no construct failed during cyclic

weight-bearing loading. Based on our study design, no conclusion regarding clinical

guidelines can be provided and further mechanical and clinical studies are warranted.

Future study design should exclude the effect from different number of screws and the

application of both locking and non-locking screws in the plate.

When preparing a fracture stabilization plan, the choices will be different based on

different intention. If one would like to provide a fixation without traumatizing excessively

the surrounded soft tissue, a MIPO technique and a longer plate working length should

be acceptable based on our study result. However, if the higher strength of the fixation

construct is the goal, we would suggest placing screw as closed to the fracture gap as

possible in each of the major fracture segment. Future mechanical and clinical studies

are needed to establish guidelines regarding the optimal application of locking

compression plates.

54

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BIOGRAPHICAL SKETCH

Peini Chao grew up in Taipei, Taiwan with her family. She graduated from Taipei

First Girl High School in 2003. She studied in the college of Veterinary Medicine in

National Taiwan University. After one-year rotational internship training in Animal

Hospital of National Taiwan University, she earned her D.V.M. in National Taiwan

University in 2009. During the summer in 2007, she completed an externship in Elwood

Animal Clinics in California.

Upon graduating in August 2009 with her D.V.M., Peini enrolled the Master of

Science program in Department of Small Animal Clinical Science in College of

Veterinary Medicine in University of Florida. She works in Comparative Orthopedic

Biomechanics laboratory, research strengthens on mechanical evaluation of plate-screw

constructs. She earned her M.S. in summer 2012.

After completion of her M.S. program, Peini will be back to Taipei, Taiwan,

continuing work in small animal practice.