Biomechanical Considerations in Osteoporotic Fracture Fixation

Abstract

Background

Osteoporotic Fractures (OF) present formidable but predictable challenges in fixation. With ageing the bone mineral density is reduced and the internal micro-architecture is disrupted. This increases fracture risk and makes implant hold tenuous. Newer implant technology has helped improve fracture fixation but the risks of early mechanical failure remain tangible.

Purpose

After fracture reduction and fixation, the surgeon remains apprehensive regarding rehabilitation. The concerns are higher in the lower limb where non-weight bearing is not possible. Understanding basic mechanics and translating that knowledge to fracture surgery helps provide secure surgical stability to enable full weight bearing assisted mobilization.

Conclusion

Applying the logic of mechanics to living biological tissue will help the surgeon better understand the unique mechanical requirements of the fractured bone. Judicious surgical technique and careful combination of implants balancing the mechanical and biological needs of the ageing broken bone will help it heal. Integrating technology and surgical technique with the established principles of osteosynthesis will help improve functional outcomes in OF.

Introduction

Human population is increasing and aging. Life expectancy is increasing and more elderly are pursuing active lifestyles with good quality of life. However, with aging, bone naturally endures physiological decline in structure and strength. Nature has intuitive adaptive mechanisms to maintain optimum bone strength with reducing mineral content. Beyond a certain threshold, the bone fragility increases and fractures occur. Weak bone is at risk of easy fracture and fragmentation. At the same time, it is weak to hold the implant too. This causes pain and loss of function depriving the individual of quality and dignity of life. Old broken living bone is programmed to heal albeit a bit slower. It only demands a better hold of the fixation device for longer duration. Newer implant technology has helped improve the fracture hold but preventing early mechanical failure remains challenging. An understanding of basic principles of mechanics and integrating that knowledge with the biology of bone healing will help improve the outcomes of managing OF.

Basic Mechanics of Bone

From a mechanical point of view, bone is a system of rigid levers, that supports and protects other body tissues and provides storage of minerals—calcium, phosphorus mainly while being metabolically a very active tissue. At the same time, bone is unceasingly busy in calcium regulation allowing bone structure to respond to stresses in real time—modeling, maintaining material properties and appropriate strength throughout life—remodeling and repair of microdamage associated with activities of daily living [1–3].

The property we as clinicians are concerned with is strength. Bone as a tissue is mechanically required to be strong, stiff, and elastic at the same time, not very heavy and yet able to absorb energy and dissipate it to avoid fracturing easily. At the microscopic level, it is explosively active contributing to calcium homeostasis, detection, and repair of microfractures, remodeling to replace older tissue and modeling in response to stresses of daily living. Osteoporotic fracture (macrolevel event) is due to an accumulation of micro- and nano-architectural impairments evolving over time and decreased bone mineral with aging [3, 4].

The mechanical behavior, i.e., strength of bone, depends on amount of bone tissue, i.e., density, the spatial distribution of the mineral content—bone geometry, and the microarchitecture of bone and tissue properties [5, 6].

This can be understood from studying the mechanical properties of bone on different scales—macrostructure, microstructure, cells and matrix. A change in any one element affects others too [7].

Bone mineral density (BMD) assessed non-invasively with dual energy X-ray absorptiometry (DEXA) is currently the gold standard to estimate bone fragility and guide the medical management of osteoporosis. The fallacy of this investigation is that it measures the bone density only. Mineral distribution within bone (bone density) contributes about 60% to the bone strength. The rest is dependent on the microarchitecture of bone [2, 5–8].

The Spatial Distribution of Bone Mineral Content—Bone Geometry

The Concept of Neutral Axis

An area in the beam/structure where the internal forces (stress) is zero on bending. Subjected to bending stresses, the outer half of the beam is under tension and the opposite half is under compression (Fig. 1) [9, 10]. An intramedullary nail lies along this axis and a plate on the tension side of bone. This places an inherently high stress on the plate from the outset compared to a nail.

Fig. 1.

Cross-section of a beam/diaphyseal bone showing the distribution of forces within the diaphyseal cortical bone on bending. The outer cortex is in unequal tension increasing toward the outer surface and the inner cortex similarly under unequal compression. In the center of the medullary canal exists a zone of no internal stress—neutral axis

Area Moment of Inertia or Second Moment of Area (Ix)

It is the geometric characteristic that measures the spatial distribution of material about the neutral axis of its structure. In physics, it is the outermost portion of the cross-section of a structure that best resists the bending stresses. Nature understands this principle and places its strongest bone quality, cortical bone at the periphery in its structure.

The area of a structure or space is given by multiplying its length with its breadth. In structure terms, the entire area of the material does not contribute equally in resisting deformation on loading. The property that determines the ability of bone or the implant to resist deformation is area moment of inertia (AMI), i.e., that portion of a structure that resists bending and represented as Ix. This is the outermost surface of bone/plate/nail on the tension side (Fig. 1).

Ix for a beam (rectangular implant—plate) = $bh3/12$ or $hb3/12$, where b is the width and h is the height of the implant in the direction of bending.

Example: A 6-inch plastic ruler held horizontal versus vertical. Held horizontally $(hb3/12)$, it can bend and break easily. Held vertical ($bh3/12)$, it becomes difficult to bend the ruler. The difference by this 90-degree change in orientation is the height. The material is distributed further from the neutral axis and that is responsible for the remarkably increased stiffness.

For a circular implant, the direction of bending is not important and the resistance to deformation depends on the fourth power of the radius (${\mathit{radius}}^4$) of the implant.

Ix for a solid cylinder = $\pi d4/64$ or $\pi r4/4$

Ix for a hollow cylinder = $(\pi (do4-di4))/64$, d_o is the outer diameter and d_i is the inner diameter.

Bone and intramedullary nails are examples of hollow cylinders.

Consider a solid pipe and a plumbing pipe made of same material and the same cross-sectional area (Fig. 2).

Fig. 2.

A solid bar and a hollow pipe made from the same amount of material. The hollow pipe has material distributed further away from the neutral axis and it is more challenging to bend the hollow pipe rather than the solid bar. Adapted from Burstein AH and Wright TM, 1994 [12]

Deformation at the outer most surface of the structure (epsilon—ε) is given by the formula

$$\epsilon =\frac{M.z}{E.Ix}$$

M is the Bending moment (load applied), z is the Distance of the material from neutral axis, E is the Modulus of elasticity of the implant material, Ix is the Area moment of inertia (distribution of material from neutral axis of the structure).

Although the area of the structures is the same (amount of material is the same), the force required to bend the hollow structure is significantly higher [10]. Bone is mechanically optimized to fulfill its role in structural stability and functions like a strong hollow pipe rather than a solid rod.

What Happens to Bone Resistance to Bending Stresses with Age?

The two possible scenarios associated with reducing bone material are:

Scenario 1: Inner diameter is same, outer diameter reduces.

Scenario 2: Outer diameter remains the same, inner diameter increases (Fig. 3).

Fig. 3.

Bone is a hollow cylinder (a). Loss of material from the endosteal side (b) versus the periosteal side (c). The AMI of bone loss from the endosteal side is higher and this is natures mechanism to maintain bone strength with reducing material with aging. Adapted from Tencer AF and Johnson KD, 1994 [1]

Applying the AMI concept, losing material on the outside is more deleterious. Resistance to bending depends on the material and its distribution (geometry). With aging, bone material is primarily lost from the endosteal surface (inner diameter). This is clinically seen in the ability to place larger diameter nails in elderly patients who otherwise appear to be small. The outer diameter remains the same, i.e., bone material having being lost and weakened is distributed further away from the neutral axis—natures inbuilt compensatory mechanism to maintain skeletal integrity to enable it to sustain function [10, 11].

The Microarchitecture of Bone

At the microscopic scale, trabecular bone and cortical bone are made up of the same constituents. The basic building block of bone is the osteon, like bricks in a building. The only difference is in their arrangement—the microarchitecture [2, 5]. At the macroscopic level, the density of osteons is lesser in trabecular bone, the packing is less parallel and more heterogeneous. Cortical bone is compact with a density of about 1.5 gm/mm³. Cancellous bone has a density less than 1 gm/mm³.

Bone microarchitecture is a system of interconnected trabeculae. The strength of bone is dependent on the (1) number of trabeculae (density of trabeculae), (2) the thickness of the trabeculae and their (3) interconnections. Trabeculae may be in the form of rods and plates. These rods and plates are interconnected to each other and are oriented along the lines of stress. Aging results in thinning of the trabeculae, increased porosity of the trabeculae, decreasing the interconnectivity of the trabeculae, and increasing the separation of the connecting trabeculae. This results in weakening of the internal structure with loss of strength (Fig. 4). Micro-CT and micro-MRI have been used to objectively assess this aspect of deterioration in bone strength [5].

Fig. 4.

Trabecular pattern—young (a) and change with aging (b) using the analogy of a bamboo ladder. With aging, the columns thin down and the interconnections between them (steps) break down. Adapted from Freidman AW. Clin Rheumat 2006 [6]

Nature has self-programmed mechanisms of compensation to counter the age-related weakening of structure. In predominantly cancellous bone, weakening is balanced by compression/impaction, increasing the strength of the bone at macrolevel, a phenomenon seen in wedging of the vertebra (Table 1). The cost of sustaining function is loss of height and pain in the back. In diaphyseal bone, loss of bone material and weakening microarchitecture is offset by widening of the medullary canal (Fig. 3).

Table 1.

What happens with age? [2, 5, 11]

Mechanics	Clinical outcome
Lesser material	Reduced bone density Reduced implant hold
Lower strength	Increased fragility
Decreased stiffness	Easily deformable
Wider canals	Thinner cortices decreasing screw hold
Increased porosity	Increased fracture risk with fragmentation Reduced implant hold

Osteoporosis is a systemic condition associated with reduced bone mineral content and disruption in the microarchitecture of bone [8]. This makes the bone fragile (break easily); brittle (fragmentation in to multiple fragments) and implant hold tenuous. The changes of aging are more pronounced in cancellous regions. In this region of long bones (juxta-articular), bicortical screw purchase is not always feasible. This makes fracture reduction and fixation more challenging. Conventional screw hold is precarious as the hold depends on implant–bone interface. The surgeon does not feel the confidence of good purchase during tightening the screw. In almost empty bone, failure is inevitable and early. On the other hand, locking screws rely on hold via their head threads tightening in the plate thread. In the case of angle stable nailing, interlocking screws depend on their threads in the center region locking in to threads in the nail hole for a more secure hold in weak bone. Multidirectional trajectories ensure higher cut out resistance to failure.

What is the Difference Between Strength and Stiffness of an Object?

Strength of an object quantifies its ability to withstand loading without failure. Failure in fracture fixation could mean either plastic deformation of the implant or break of the implant. This is a property of the structure of the implant, i.e., a combination of its material and geometry.

Stiffness of an object defines how difficult it is to deform the material under loading. This is a property of the material exclusively and is quantified by the Young’s modulus (modulus of elasticity) [9].

Mechanics of Implants

Metals enjoy wide application in orthopedics as structural load-bearing devices, partial and total joint replacement devices, instruments, splints, braces, and traction apparatus. The reasons for this popularity are:

1. Metals possess a high modulus of elasticity with reasonable yield strength that will bear significant loads without large elastic deformation or permanent deformity.

2. Metals possess a high endurance limit (fatigue strength) that allows cyclic loading without fatigue fracture.

3. Metals have high ductility and ultimate tensile strength. Stresses exceeding the yield point will produce plastic deformation rather than sudden fractures, permitting measures to rectify the fault.

4. Metals can be fabricated into a variety of parts by conventional techniques. Alloying of metals allows adjustment of mechanical wear and corrosion properties.

5. With care in surface finishing, metals have good resistance to environments encountered in vivo [1, 9, 12].

Metal Structure

The properties and performance of metals and their alloys is dependent on the composition and structure of the material. How an alloy is fabricated will affect the structure, which in turn will dictate the mechanical properties. The usual route to the development of an alloy of commercial value is to select a base element and then add small amounts of other elements to change and enhance the properties of the base metal.

Metals are made of crystals joined together at grain boundaries, much like individual osteons in the bone with its boundaries (cement lines). Physical and mechanical properties of metals are governed by relationship between atoms within a crystal and between crystals. These crystals are not defect free and defects in fact are the primary factors affecting properties of an alloy. The defects possible are:

1. Point defects: These occur when a lattice site in the crystal is not occupied by an atom. These defects are present in all metals and alloys and provide a mechanism for diffusion in solids.

2. Area defects: These are the grain boundaries, i.e., the region where two crystals come together.

3. Line defects (dislocations): These are the major defect in manufacturing that affect the mechanical properties of metals. These are the result of an extra half plane of atoms in a crystal that locally distort the crystal structure. When stress of sufficient magnitude is applied, these dislocations move within the lattice resulting in permanent change in shape, i.e., plastic deformations.

4. Volume defects: These are voids or cracks. Small scratches on the surface of a metal can significantly lower the corrosion resistance and fatigue strength of the alloy. Titanium is extremely notch sensitive.

Processes that tend to impede the motion of dislocations fill up the point defect and reduce the area and volume defects that will act together to strengthen the alloy. Evolution of an alloy from its elemental ore form to the finished product incorporates processes to provide optimum quality implants.

Evolution of Metallic Component

Most of the metals in nature occur in impure form (ore) mixed with oxygen and elements such as nitrogen, silicon, carbon, sulfur, and other metals. Each component of the alloy is first extracted from the primary source, reduced to the metallic form, and cleansed of unwanted impurities. The alloy is then fabricated by combing appropriate elements and melting them under suitable conditions. This process incorporates a strengthening mechanism known as solid solute strengthening, i.e., addition of one or more elements (solute) to the base metal (solvent). Point defects in the metal base are reduced. Dislocations are pinned down by developing locally solute rich regions in the vicinity of the dislocation line and strength is increased significantly. For example, addition of < 0.1% of carbon to iron results in steel with tenfold increase in strength. Addition of small amount of oxygen to titanium significantly increases yield and fatigue strength.

Mechanical Properties of Material and Structure [1, 4, 9, 12]

Familiarity with some of the terminology used by material scientists to describe the strength characteristics of metals is necessary to make an informed selection of implants. Some of the terms used are:

Elastic limit: The maximum load that a metal can withstand and yet assume its original shape when the load is discontinued.

Yield strength: The stress beyond the elastic limit that results in permanent deformation.

Modulus of elasticity: This is a measure of stiffness. It is the proportionality constant in the linear portion of the stress–strain curve below the yield point. It is an intrinsic property of the material generated by attraction of atoms within the material and it has essentially no variation with thermal and mechanical history. A high modulus of elasticity indicates that the material is stiff, like ceramic.

Fatigue strength (endurance limit): This is the maximum load that a metal can withstand without fracturing when subjected to 10 million cyclic loads. This is the most important characteristic in a femur stem of the total hip replacement (THR) and can be significantly altered by heat treatment.

Ultimate tensile strength: The maximum load that a metal can tolerate in a single application without breaking. If this is exceeded, the metal will fracture.

Fatigue limit: The load that a metal can endure indefinitely without bending or breaking when subjected to loading.

Fatigue fracture: The break of an implant caused by repetitive application of loads in the range between fatigue limit and yield strength. Repetitive loading below the yield point (within the elastic limit) gradually degrades the mechanical properties over time. This is called fatigue. Why does this happen? This occurs due to formation of small cracks. With time, the cracks propagate till ultimate failure. Like metals, the fatigue strength of bone too is significantly lower than static strength (as calculated in laboratory conditions). Bone is an excellent fatigue-resistant material because of its ability to repair itself (remodeling) and its ability to resist crack propagation. As a tissue, bone is not robust in preventing crack initiation. Microfractures are inevitable and associated with activities of daily living. These microfractures tend to spread. Natures’ inbuilt defense mechanisms arrest this crack spread by exiting at cement lines—the natural boundaries between osteons. Cancellous bone with its increased porosity arrests the spread more efficiently than the densely packed cortical bone. Increasing porosity with aging is a faint but mechanically logical attempt to slow crack propagation in osteoporotic bone [4, 10].

The mechanical properties of all alloys are sensitive to differences in processing. Although the elastic modulus is an intrinsic property of the metal, yield strength, ultimate tensile strength and fatigue properties are strongly affected by small changes in chemical composition and even more so by the phase structure created during fabrication.

Stainless Steel (SS)

Historically, steels were the first modern metal alloys to be used in orthopedics. These are iron-based alloys with low carbon content and high chromium content. The controlled presence of carbon is what distinguishes steel from other iron alloys. Chromium permits the formation of a stable chromium-oxide-containing surface layer during passivation. Low carbon content controls the formation of carbides, which are needed for surface hardness and strength. The downside is that it increases in vivo corrosion. Molybdenum is added to control corrosion but is expensive and less effective than chromium. It hardens the alloy making it difficult to be worked. Other elements are added to control mechanical properties.

Stainless steel is easily worked and forged. In the annealed state, it can be fabricated by virtually all finishing processes. This versatility combined with the low price of the elemental components contributes to the wide popularity in orthopedic applications. Stainless steel used in the early days had inadequate yield and fatigue strength with unacceptable incidence of fatigue failure. However, newer forged alloys with nitrogen added have shown exceptional mechanical properties and stainless steels are poised to make a revival in orthopedic implants.

Titanium (Ti)

Titanium-based alloys have entered commercial use much more recently than SS and cobalt-based alloys but are the most intensely studied alloys systems for orthopedic applications. They are well known for their superior tissue compatibility. Commercial pure titanium is not as strong as its alloys. Small amounts of gaseous impurities specifically oxygen and carbon add to the strength. Ti-6Al-4V is the primary implant alloy in use today. The modulus of elasticity is half that of SS and cobalt-based alloys. This reduced modulus is thought to aid in stress transfer from implant to the bone, thereby minimizing the potential for stress shielding and bone resorption. It has high fatigue and yield strength. Recent reports regarding the toxic effects of Vanadium have stimulated research for newer alloys. Ti-5Al-25Fe is currently under research. One of the less desirable effects of titanium alloy is their notch sensitivity. This is an intrinsic property that cannot be changed by working or processing. It is, therefore, even more important for the alloy systems to avoid scratching the surface.

Titanium alloys have two major phases, an alpha phase that is table at high temperature and beta phase that is stable at low temperature. Addition of aluminum tends to stabilize the alpha phase, whereas vanadium stabilizes the beta phase. This is why it is described as an alpha + beta alloy. The presence of these two phases permits generation of diverse microstructures with different mechanical properties by heat treatment or chemical addition.

Mechanobiology of Fracture Healing

The study of how the mechanical environment of the fracture modulates biological response of the fracture healing process [13]. The variables are (Table 2).

Table 2.

Variables of how mechanics can modulate biological responses [1, 4, 12]

1	The material of the implant
2	The structure (geometry) of the implant
3	The fracture patterns
4	Bone quality
5	How the implant is fixed to bone?
6	Loading by the patient—static and cycling (repetitive)

The Challenge—Balancing Strength and Stiffness of the Fixation

Knowledge of principles is good. An understanding will help in correct surgical choices and techniques.

1. Material: Independent of shape and size, study of material in laboratory allows comparison of the behavior of two or more solids to loading/deformation. The properties studied are stress, strain, and elastic modulus. The tested material is preset to a fixed dimension. The testing is usually performed in tension (stretching). The other pure deforming force is compression. The other fixed variable is that static loads are studied. However, in nature, bone and implants are fundamentally subjected to composite stresses—mixed, cyclical, and repetitive of which the most studied and understood are bending and torsion. Bending induces simultaneous tensile and compressive forces on opposite sides of the material (Fig. 1).

Elastic modulus (Young’s modulus) is the most important property that helps compare the clinically relevant behavior of materials. It is the relation of the stress to the strain of a material on stretching. All materials are assumed to have a linear relation till their yield point (Fig. 5). SS has a modulus of elasticity of 200 GPa, almost double of Ti alloys (100 GPa). For a given load, a Ti alloy implant will deform more than SS implant of same dimension.

Fig. 5.

Relation of stress to strain for a material under tensile stress. Up to the yield stress, the material behaves elastically. On removal of stress, the original shape is restored. Beyond the yield point, the material deforms permanently. On removal of stress, the material will not regain its original shape. Adapted from van der Meulen MCH, et al. Orthop Knowledge Online J 2016 [4]

The behavior of the material after the yield point determines whether the material is brittle or ductile. No deformation makes the material brittle (ceramic).

The other material property used to compare material behavior is ultimate tensile strength (UTS)—the load at which the material breaks catastrophically. The UTS for SS is 1300 MPa and Ti alloy is 950 MPa.

2. Structural property of the implant: The most useful and clinically relevant property to be understood. Testing of an implant on loading produces a load-deformation graph like material testing but the terminology used is slightly different (Fig. 6).

Fig. 6.

Relation of stress to the strain for a structure (material with its geometric shape). The shape of graph is similar; only the terminology used is different. Stress = load, strain = deformation. Adapted from Gordon JE, 1978 [10]

Structural property depends on the material and structure. The property relevant to geometry is area moment of inertia (AMI). Clinically, the effect of material stiffness is less relevant compared to structural stiffness. AMI can be influenced by (1) choice of the implant, (2) reduction of the fracture and (3) placement of the implant—altering the height of the implant changes the stiffness to deformation by the third power to the height of the implant [1, 4, 17].

When the fracture is simple and can be accurately reduced (intrinsically stable), the implant and bone both contribute to resisting bending forces.

When the fracture is multifragmentary and the opposing fragments cannot be approximated, the AMI is solely dependent on the implant. Balancing the strength and stiffness then becomes the challenge. Structural properties of the implant and bone–implant construct determine the mechanical environment of fracture healing [18, 19]. The structural stiffness varies with the third power of the height of the plate in the direction of bending. This variable can be surgically controlled. For example, a plate on the posterior surface of elbow versus the plate on the lateral surface (Fig. 7).

Fig. 7.

The concept of AMI. The plate on the posterior cortex of distal humerus resists deformation by its thickness—2 mm. On the other hand, the medial plate applied on the medial column of distal humerus resists deformation by its width—6 mm. The mechanical increase to deformation is by a factor of 28

3. Working length: Span of the plate across the fracture. The larger the span, the lower is the stiffness of the implant–bone construct, more the deformation of the implant and higher the strain on healing tissue (Fig. 8) [14, 15].

Fig. 8.

Working length of a plate. For a fracture with gap (a) construct (b) utilizes two screws to hold the fracture. The distance between the inner two screws determines the working length. Construct (c) also utilizes two screws with the inner two screws at a distance further from the fracture. The working length marked by the red line is longer. Longer working length allows for more elasticity in the construct. On loading, if the opposite cortices touch each other, (d) the bone shares load with the implant reducing the risk of implant failure

The working length must be enough to allow deformation of the plate/nail within its elastic limit so that on loading, the tissue strain in the fracture zone is < 10%. Over time, callus is desirable. It indicates viable bone actively participating in union and sharing load with the implant. This is the true marriage of an implant with a broken bone. Healing tissue strain depends on implant stiffness (material property) and span (working length) [16].

1. When opposite cortices touch each other, bone contact shares the bending forces, thereby protecting the plate. This is possible in a simple fracture pattern (Fig. 9).

2. If the fracture is multifragmentary and the opposite cortices cannot touch each other, the fixation relies on intact biology to protect it from failure. Callus on the side opposite to the plate over time neutralizes the bending stresses.

Fig. 9.

The mechanical advantage of a nail over a plate applied in the distal femur. The nail is along the anatomical axis (blue strip line) and closer to the mechanical axis (black stripe line). The plate on the other hand is placed on the tension cortex and at a distance from the mechanical axis. This imposes higher bending loading on the plate (red arcs) compared to the nail (blue arcs). The strain around a nail is uniform on healing tissue (blue waveform) and asymmetric after plating (red wave form)

Difference between a long plate and short plate for a fixed working length and similar fracture pattern: A longer plate allows for greater spacing of the screws. This enhances the resistance of the fixation construct to repetitive translational (axial) deformation and torsional stresses. Overall, the fatigue life of the construct is better. Fracture healing, however, depends on the span (working length) and will remain unchanged.

The strength of the implant + bone construct allows mobilization: This is important in the initial phase of surgery to enable pain free rehabilitation of the nearby joints. Load transmission is permitted by the implant across the fracture. Leverage in the fragments is crucial (hold). The second factor in anchorage is durability—ability to hold on till bone takes over its mechanical job. This depends on implant strength and its hold in bone. In conventional screws, implant hold depends on screw purchase in bone. This purchase is largely achieved in cortical bone. With thinning cortices in osteoporosis, the hold becomes weaker than healthy bone. In juxta-articular regions, bicortical screw purchase may not always be feasible. Enhanced thread depth of the 6.5 mm cancellous screws (1.75 mm versus 0.75 mm for 4.5 mm cortical screw) marginally improves the hold but the durability remains low. The anchorage of conventional screws in this region is insecure in old bone. Locking screws provide a more secure hold because their anchorage relies on the screw head locking in the plate [14, 15]. Multidirectional trajectories increase the pull-out resistance enhancing the durability.

The stiffness of the implant + bone construct facilitates healing: This depends on the span (working length). This variable is vital for the formation of bone tissue over time. This determines the deflection of the implant that permits favorable deformation (strain) of healing bone tissue [16, 17]. The strain must remain under 10% gradually reducing over time to less than 2%.

Deformation for the same AMI is related to the third power of span of the implant.

The formula is $\frac{\text{Load}\times {\text{length}}^3}{\text{Modulus of Elasticity}\times \text{AMI}}$

For a load of 300 N and a span (working length) of 2 cm, the deflection of a SS 3.5 mm LCP (MOE—200 GPa, AMI—40 mm⁴) implant is 0.3 mm. When the span is doubled to 4 cm, the implant deflection increases to 2.4 mm. For a 3.5 mm LCP made from Ti alloy (MOE—100 GPa, AMI—40 mm⁴), the deflection will be 0.6 mm and 4.8 mm, respectively. Excessive implant deflection puts higher strain on healing tissue and may interfere in progression of maturation of healing callus. On the other hand, too little deflection (deformation) may not initiate the process of hematoma differentiation.

Fracture Healing from the Healing Tissues Perspective

What does healing tissue need? It needs blood and some movement [14, 15].

Where there is no gap, blood vessels can cross over directly. This is intramembranous ossification, i.e., the formation of bone without intermediate cartilage formation. This is desirable in intra-articular fractures.

Fracture reduction remains paramount and the first step in surgical reconstruction of the broken bone. Bone, even when porotic, remains stronger in compression than tension. End to end apposition though desirable is not preferred in the porotic bone. A certain degree of cortical overlap is ideal (impaction). When fracture ends contact each other across the fracture, the implant is essentially in a load-sharing mode with the bone.

Difference Between a Simple and Multifragmentary Fracture

Simple fracture: The injuring force is lower and soft tissue disruption around the fracture site lesser compared to multifragmentary fractures. This is essentially a biological property of the limb.

In a multifragmentary fracture, from a mechanical point of view, the broken ends of the bone cannot be made to touch each other after reduction. The implants are essentially load-bearing and the strength and durability of fixation depend on the purchase of implants in bone. The nail scores over plates due to their intramedullary location. Why is this? [1, 12, 20].

1. The concept of neutral axis. No bone in the human body is straight. Being curved in the coronal and sagittal planes, on standing or functional loading, one side of the bone is under compression and the opposite cortex will be in tension (Fig. 1). Taking femur as an example, in the coronal plane, the medial cortex is under compression and the outer cortex in tension. Somewhere within the medullary canal, the stresses convert from compression to tension. There must be a line where the bone is neither in tension nor compression. This is the neutral axis of the bone (Fig. 1). An intramedullary nail lies in this axis and on functional loading, causes a uniform deformation of healing tissue all around it (symmetrical strain). On the other hand, plates are applied on the tension side. On loading, plate deformation leads to asymmetrical tissue distortion. This distortion is least under the plate and increases toward the opposite cortex (asymmetrical strain) (Fig. 9). The clinical correlate is very less bone formation under the plate (near cortex) and more callus on the opposite cortex of bone (far cortex). In the sagittal plane, the anterior cortex is in tension and the posterior cortex in compression. Peri-implant fractures essentially begin on the tension side and then progress across the bone diameter. Once a critical threshold is reached, catastrophic failure happens. Clinical correlate is the atypical femur fracture on the lateral cortex and peri-implant fracture through the most proximal locking screw starting from the anterior cortex in distal femur plating using longer implants.

2. Bending loads on the plate. The plates are applied to the tension cortex of bone. In the femur, the plate is lateral from the neutral axis of bone and the mechanical axis of the limb. This places the plates at a mechanical disadvantage and they are subjected to predominant bending stresses (Fig. 10). In real life, functional loading after fracture fixation is repetitive, i.e., cyclical.

Fig. 10.

(Section 1): Locked screws function as a team. Conventional screws rely on friction between plate undersurface and bone (physical contact) to counter stress (a). On loosening, conventional screws toggle individually eroding bone (b). Locking screws move together and eat away more bone volume prior to failure (c). On pull-out, conventional screws move along their axis of insertion (d). Locking screws with their head fixed to the plate erode greater bone volume before failure (e). Adapted from Cornier P, et al. OTSR 2010 [19]. Clinical example of femur diaphysis bone eaten out by locking plate failure with screws locked to the plate in the distal fragment (f) and the proximal fragment (g)

Osteoporotic bone too will heal if given time. Formation of callus on the side opposite the plate helps share the load and reduces the chance of failure. In this race of bone union versus implant failure, the vitality of the broken bone and its surrounding soft tissues is critical. The final post-operative X-ray may not inform us about the finesse of the surgical technique. Only time tells.

The strategies to address the issues of stability to enable fracture healing and immediate post-operative mobility can be achieved by a combination of selection of the implant and judicious surgical technique [20, 21].

1. Implant: Use of locked plates [22, 23].

2. Surgical technique: (1) Avoid absolute stability, (2) avoid placement of unicortical screws, (3) combination of implants, (4) filling up voids/bone defects, and (5) augmentation of the implant to enhance its hold [24–26].

Locking Screw Plates

Locked plating offers two distinct advantages over conventional plating. First, it creates a fixed angle construct and second, all the locked screws on one side of the fracture function as a team [18, 19, 22, 23]. A fixed angle construct has the following advantages:

1. A fracture well reduced by the surgeon is maintained well with lower risk of screw loosening. The hold of the screw is primarily by locking in the plate and less dependent on the quality of bone.

2. Bicortical purchase though desirable is not essential for more secure hold. This benefit is seen in periarticular fractures where bicortical purchase is not possible and the screws in the articular block are long but unicortical.

3. On functional loading, the stress across the fracture is transmitted through the plate and screws. Based on the surgical technique adopted, on loading, locked screws and plate deform rather than the injured bone. When this deformation of the plate is within the strain limit of healing bone tissue, the fracture will heal.

4. Locked screws on either side of the fracture function together as a team. On the other hand, conventional screws act independent of each other. Locked screws are mechanically constrained to function together because of the screw head locking into the plate hole. If the implant–bone composite becomes unstable, all the locking screws on one side of the fracture will deform together. The volume of bone eroded by all the failing screws together is much larger than conventional screws failing independently (Fig. 10). This mechanical advantage gives locked plate technology a significant mechanical advantage especially in osteoporotic fracture fixation.

Taking all factors in account, screws locked in a plate give healing bone more time to unite. Locked screws succeed and fail as a team.

Avoiding Absolute Stability in Meta-diaphyseal Fractures

The preconditions for achieving absolute stability are (1) anatomical reduction, (2) absence of fragmentation, (3) good bone quality, and (4) compression in a plane perpendicular to the fracture.

Osteoporotic bone is fragile with thinned out cortices. Hidden fragmentation is the rule rather than the exception. Anatomical reduction may be achievable but attempts to compress the fracture may fail. Compression relies on good hold of the screw in the thin far cortex and good resistance of the near cortex to prevent the head of the screw from sinking into bone. What is desirable is impaction to enable good bone contact across as much of bone perimeter as possible across the fracture so that the implant is in a load-sharing mode with the broken bone.

Combination of Implants

Intramedullary implants lie in the neutral axis of long bones and are mechanically superior to plates applied on the tension surface of bone in load transmission across a broken bone. When the fragment size is small and locking screw hold possibly tenuous in one fragment, it is mechanically logical to add a lower profile plate to enhance fracture stability. The plate adds another column of support and the screws through the plate increase stability by additionally working as Poller devices constraining the nail in the wide medullary canal (Fig. 11).

Fig. 11.

Combination of implants to maximize stability to enable early function. Long spiral multifragmentary distal femur fracture (a), stabilized by distal femur nail + distal tibia plate construct (b) and 6-month follow-up (c) showing healing by secondary intention

Orthogonal plating with one small reduction plate and another large protection plate too can enable good stability and early full weight mobilization. The reduction plate is applied by open reduction and the protection plate placed by minimally invasive method spanning as much length of bone as possible. The protection plate must be contoured and no conventional screw must be used. Orthogonal plating significantly enhances fracture stability and prudent surgical technique will avoid biological damage to endear satisfactory fracture healing (Fig. 12).

Fig. 12.

Combination of implants to avoid acute fixation failure in a severely osteoporotic bone. Midshaft transverse widely displaced humerus fracture with proximal humerus nonunion (a), stabilized by a short 3.5 mm reduction plate (open reduction) and long 4.5 mm protection plate (minimally invasive) AP view (b) and lateral view (c) with union at 6 months AP view (d) and lateral view (e)

Filling Up of Voids and Defects

Aging bone is soft and metaphyseal bone occasionally crumples into itself. Restoration of axis and alignment are mandatory but some impaction is always desirable. Impaction enables cortical bone contact along the perimeter of bone. This is good for implant stability. Presence of voids within the bone (cancellous area) make it difficult for healing cells to jumps across large gaps. Filling the voids with bone graft (allograft or bone substitutes) increases structural stability of the composite of bone and metal. Cement fill is more complete and compact compared to granular bone graft. Cement also enhances the holding power of the screws and significantly reduces the risk of secondary displacement. The osteoconductive scaffold provided enables cells to migrate across the voids better for bone ingrowth.

Augmentation of the Implant

The greatest risk of surgical fixation of an osteoporotic fracture is quality of implant hold. Most fixation devices are screwed in and some hammered in (spiral blade). The tactile feedback of a conventional screw tightening is reassuring to the surgeon. The spiral blade was theoretically thought to remove no bone but compact the fragile bone for superior hold. Without tactile feedback, it was designed to provide good hold. Unfortunately, aging bone is devoid of material to hold or compact and acute fixation failure is possible (Fig. 13). Augmentation of the screw/blade device with cement substantially increases implant purchase in cancellous bone, enabling good pain relief early mobilization and decreased risk of secondary loss of reduction (Fig. 14).

Fig. 13.

The tenuous hold of a spiral blade in a 91-year-old lady with intertrochanteric fracture. A simple intertrochanteric fracture (a) showing good reduction with the location of the spiral blade deep and in the center of the head (b). The spiral blade fell out within a week of surgery (c)

Fig. 14.

Cement augmentation of the lag screw substantially increases device hold reducing the risk of loosening

Conclusion

Osteoporotic fractures are challenging to manage medically and surgically. The elderly bone is programmed to heal albeit a bit slower. The principles of fracture fixation remain the same as healthy young bone but demand more careful adherence. An understanding of mechanics will help the surgeon better manage these injuries. Judicious surgical technique, thoughtful combination of implants balancing the mechanical and biological needs of the aging broken bone will help surgeons deliver better care to this special group of patients. With current knowledge and instrumentation at our disposal, we can give good clinical results to our patients; permitting early full weight-bearing-assisted mobilization, minimizing acute fixation failure, and allowing the broken bone to heal.

Declarations

Conflict of interest

The author declares no conflict of interest. The images included in the manuscript are all original. Sketches have been made by the author.

Ethical standard statement

This article does not contain any studies with human or animal subjects performed by the any of the authors.

Informed consent

For this type of study informed consent is not required.