- THIS MATERIAL IS PUBLISHED AND PROTECTED BY U.S. COPYRIGHT LAW - REPRODUCTION PROHIBITED UNLESS FOR PERSONAL USE, EXCEPTING AUTHOR PERMISSION - ORTHOPEDIC BIOMATERIALS: THE ORTHOPEDIC SCREW Peter F. Kelly, D.P.M., F.A.C.F.A.S. Diplomate, American Board of Podiatric Surgery Fellow, American College of Foot and Ankle Surgeons ABSTRACT An overview of the design characteristics of the metal biomaterials is presented, and the orthopedic screw is used as an example. The major engineering aspects are considered, so that the reader may have the opportunity to employ these foundations knowledgeably in the future evaluation of other orthopedic metal materials when they become available. The Significance of Having Biomaterials The design of orthopedic biomaterials is generally a simulation of nature by man and his technology. For temporary reconstructive procedures uncritical geometries of a highly variable nature may be used. On the other hand implants, prosthesis, and fixating elements must additionally function within the rigid requirements of natural materials. Both must be satisfactory for emulation of patient comfort, minimal tissue response, and have long-term mechanical stability. Considerations for Design 1. Design for Rigidity and Strength It is understood that biomaterials must withstand the loads to which they are exposed without distorting, deforming, deteriorating or breaking. Rigidity in orthopedic applications is as equal in importance as the strength of the device. Strength may be viewed as the ability to reconform to its original shape after a deformational load is released. In physics the definition of this is elasticity. Rigidity is the amount of deformation placed upon the device up to the point of the elastic limit. The elastic limit is determined by 1. the magnitude of the force, 2. the cross-sectional area, and 3. Young's modulus, in the ideal model. 2. Designs to Counteract Corrosion The use of alloyed metals greatly increases the strength of the crystal lattice. This is necessary for applications of highly ductile materials which are easily shaped to the bone and combined with hard, wear resistant materials used for joint surfaces. Also the surface approximation of screw and plate fixating devices each contact dissimilar portions of their opposed crystalloid material. Crevice corrosion which results will be discussed shortly. Combining two or more dissimilar metals provokes a galvanic current in an electrolytic media. Protection of one metal ("noble" metal, cathode) from dissolution by corrosion of the other ("base" or "reactive" metal, anode) is the nature of the problem. Metals with as similar an electromotive force (EMF) potential as possible should thus be used. A film generated by the reduction of oxygen by this micro-current will be shown to cover these materials in vivo, and thus the alloys, if similar enough, may react somewhat less directly with the electrolytes in the environment. Using a table of Normal Potentials provides a rapid screen of metal combinations. Anodic protection: Generation of passive film, or plating by inert materials will maintain passivity of reactivity, but external forces may provoke damage to the material so that production passivity may not be accomplished. The goal is an alloy combination which requires very little current for passive film formation so that anodic protection might be both produced and maintained. Ruthenium has been shown to be the most effective noble metal resisting chloride breakdown and active dissolution. This application of corrosion control is useful in surgical alloys where the rate of dissolution is minimal between anodic and cathodic metals and is not increased by thin coupling. Titanium-palladium alloys are a demonstration of this application. Cathodic protection: A second metal is connected to the metal of interest, with the goal of lowering the corrosion driving potential. This might be highly applicable to orthopedics were it not for the second metal's "auxiliary anodic dissolution" product accumulating with an unknown deleterious toxic response. Implanted batteries, with neutral and electron generating electrodes used to protect the specimen, might prove to be a cumbersome alternative. Thus cathodic protection is not considered in orthopedic biomaterials. Certain passive alloys: Unique characteristics of thermodynamic resistance are shown when passive alloys are immersed in body fluids. Crevice corrosion which results from low EMF activity dissolution, and pitting from high activity potentials can be avoided if the metal concerned has a potential which is less electronegative than the oxygen reduction equivalent and more electronegative ("passivitation potential") than water (or hydrogen ion). 3. Corrosion and Cell Function With regard to corrosion and cell function, none of the metals used in current orthopedic applications produce deleterious responses in cell metabolism, matrix damage, chemical environment changes (production of hydrogen or hydroxyl ions, or gas evolution), or cell migratory behavior. These would be the major expected alterations potentially affecting osteogenesis, collagen fiber allignment, and precursor cell transformation, to name a few. 4. Metallurgic Design The highly electrolytic, corrosive environment of the body, in addition to the cellular reactions of metallic dissolution, narrows the selection of materials to only those based on the noble metals, titanium, zirconium, iron, cobalt, nickel, tantalum, silver, and gold. Tantalum and the noble metals exclude themselves from construction by their unsuitable physical properties, and zirconium is too expensive. Unfortunately at present, titanium alloys and probably zirconium, niobium, and tantalum based materials are the only ones which fulfill the passivitation potential requirement of the chloride media. Moreover, these alloys may be intercombined and will continue to demonstrate passive behavior. Lattice ion positions or crevice appositions show unchanged rates of dissolution when any of these constituents are used. Unique problems are imposed by crevice corrosion. A passive metal in this case will behave as two anodes occurring in parallel. Throughout the majority of the metal surface, and through the passive oxide film, there is a slow rate of dissolution and reduction of oxygen. However at the crevice there is a rapid, active dissolution. Combinations with a second passive metal will effectively enlarge this metal's surface and will serve to reduce the rapidity of the dissolution in the otherwise more concentrated and oxidized area. The junction between a screw and a plate may suffer mechanical abrasion and thus remove the protective oxide film. Here the same principles apply. As an indication that the surface was filmed over the electrode potential should return to the level consistent with both alloys. However formation of the oxide film may be inhibited by the geometry of the crevice. Alloy composition influences lattice structure, and therefore physical and mechanical properties. The surface contacting tissue is desired to be as polished and biologically, as well as electrochemically, nonreactive as possible. Therefore metals similar in size in their atomic radii which promote continuous solubility of each other over the lattice, will contribute to the stable physical allignments necessary. Crystal structures which more closely approximate each other influence solubility and homogeneity of the material. The significance of this is in formation of intermediate phases which diminish strength of the material, smoothness of metal surface, and resistance of the surface to abrasion. For example, nonmetals (carbon and nitrogen) must be present in low concentrations due to their significantly smaller atomic radii. 5. Discussion of Materials Iron has a variety of alloys but only the stainless steels are applicable in orthopedic surgery. Additions to ferric alloys will be considered in the next few paragraphs, since they are currently the most popularly used. Brief mention will be made of other alloys. Carbon: In many steel alloys, carbon has a greater affinity for the alloying elements in the stainless steels. For example with chromium, a chromium carbide is formed to the exclusion of both of their participations in the lattice. Depletion of chromium below 12% in steel alloys decreases its passivity and corrosion resistance markedly. Chromium: Along with iron, chromium is a participant in the formation of a film which possesses corrosion resistance. Above 13% chromium concentration, ferrite is the stable phase at all temperatures of cooling during alloy manufacture. This means it retains the body centered cubic (BCC) structure characteristic of the (ferrite) and (iron) phases, and not the face centered cubic (FCC) structure of the (austentite) phase. (See Fig. 3 and 4). Nickel: Should nickel at 3% be added to 13% chromium steel and the alloy quenched rather than slowly cooled, a metastable phase (martesite) is formed, having a strained body centered tetragonal structure. These are the strong stainless steels used for cutting blades, but do not have the high passivity of the austenitic steels. Therefore in the manufacture of orthopedic bone screws the thread strength is maintained by a partial austenite to martensite conversion. The austenitic stainless steels are the most corrosion resistant. Nickel expands this austenite range and allows larger amounts of chromium and molybdenum to be used, both of which add strength and corrosion resistance in acid and chloride solutions. Alloys used in orthopedic surgery are AISI (American Iron and Steel Institute) type 316 and 317. THE ORTHOPEDIC SCREW Goals 1. Load transfer and fatigue protection: reduction of unit stresses, distribution of torque along thread plane, and equal distribution throughout other screws if plates are used. Extraction loading is dependent upon: 1) screw thread area, 2) thread number (in cortex particularly), and 3) the outside, or major diameter. Shearing loads are due to perpendicular sliding of the bone cortex, or of the plate near screw head. Flexural loads increase stresses by the fourth power of the diameter, (d2 - d1)4, so a reduction by one half the minor diameter produces 16 times an increase in the flexural load. Tortional screw strength is related to the cube of the minor diameter and ultimate tensile strength related to the square of the minor diameter. Therefore it is important to realize that deeply cut threads, while increasing load transfer, will at some point offer an optional load transfer (screw thread to bone) but only in a balance with intrinsic elasticity in response to deformational forces. This does not include aspects of safety factors of intrinsic screw force distributions. 2. Minimal trauma of insertion and removal: as the procedure involves stripping the blood carrying soft tissue from the periosteal bone. 3. Balance between 1. and 2. above: a trade-off between strength and good healing potential. 4. Role of screw head: transformation of shearing forces to axial forces is accomplished by the triangular wedging recessed in the conical distal surface of the screw head. This acts as a wedge forcing the screw to pull on its threads each time the bone cortex or plate works sideways. This also reduces the prestresses of perfect alignments which can cause fatigue leading to screw failure. 5. Role of adequate purchase: to maintain fixation, the screw must be anchored in a secure bony cortex. Axial forces will easily strip cancellous bone. Cancellous bone is composed of only 20% trabecular bone and 80% soft tissue. Cast immobilization should thus be used should screw fixation be used in cancellous bone structures. SCREW STYLES Four major styles are used: 1. Cortical - self tapping cross slotted head 2. Cortical - non self tapping hexagonal socket 3. Cancellous hexagonal recess 4. Malleolar - self tapping hexagonal recess (Designs 2, 3, and 4 are AO group developed) 4. Malleolar - self tapping hexagonal recess (Designs 2, 3, and 4 are AO group developed) FIGURE 1. "Schematic diagrams reveal the types of loading which are imposed upon surgical implants." (Taken from Materials and Orthopaedic Surgery, Mears, D.C.) FIGURE 2. "Load-deformation curve for a beam made of a ductile material. The elastic range of the curve is between a and b. At point b, some of the material reaches its elastic limit. At point c, all of the material at a particular section has reached its elastic limit. At point d, the beam ruptured." (Taken from Orthopaedic Biomechanics, Frankel, V.H. and Burstein, A.H.) FIGURE 3. "A. The constitutional diagram for iron is shown. As molten iron cools, initially, it freezes or solidifies as iron or ferrite, with a body- centered cubic (BCC) structure. On further cooling it converts first to iron, or austenite, with a face-centered cubic (FCC) structure; then it converts back to BCC, known as iron." FIGURE 3. "B. The effect of carbon on the constitution of iron is shown. The addition of small quantities of carbon depresses the freezing point of molten iron, elevates the temperature at which BCC converts to FCC and depresses the temperature at which the alloy reverts to the BCC structure. It thereby widens the austenitic range." (Taken from Materials and Orthopaedic Surgery, Mears, D.C.) FIGURE 4. "A. In the closed loop of austenite, the influence of supplementary chromium of the constitution of low carbon steel is shown. With elevation of chromium concentration the temperature at which BCC transforms to FCC is elevated and then lowered. Conversely with a diminishing concentration of chromium conversion back to BCC occurs initially at a lower temperature but subsequently at an elevated temperature. In this way the austenite range forms a closed loop." FIGURE 4. "B. The effect of supplementary nickel to a chrome steel with concentration of chromium exceeding the range at which the closed austenitic loop prevails is shown. With higher concentrations of nickel, the austenitic structure FCC reverts to a stable phase at room temperature. There is an intermediate range, however, when cooling the austenitic structure partially transforms it into ferrite and forms the hard martensitic steels." (Taken from Materials and Orthopaedic Surgery, Mears, D.C.) FIGURE 5. "Schematic views and photographs show three types of screws: A, 4-mm self-tapping screw; B, 4.5-mm AO cortical screw; C, 6.5-mm cancellous screw; and D, a self-tapping malleolar screw." (Taken from Materials and Orthopaedic Surgery, Mears, D.C.) TABLE 1. -- ORTHOPEDIC STEELS Orthopedic Steels Mb Cr Ni AISI - 316L 2.5% max. 18-20% 10-14% AISI - 317 3.5% max. 18-20% 10-14% Note: 1. combined minimum > 24% 2. Cr > 8% minimum 3. Ni > 8% minimum TABLE 2. -- OTHER ALLOYS Name Constituents Corrosion Wear Comments Inconel Nickel - Chromium High High (75%) Nimonic Nickel - Chromium - Iron High Good Creep (20%) (5%) resistance Nickel - Gold High High Expensive Nitinol Nickel - Titanium High High "Memory"- after plastic deformation MP 35 N Nickel - Cobalt - Cr - Mb Excellent Excellent Expensive (Protasul-10)(35%) (35%) (20%) (10%) Stellite 21 Cobalt - Chromium - Mb - Ni Good High High costs (61%) (28%) (6%) (2%) of quality control for each item Stellite 25 Cobalt - Cr - Ni - Tungsten Good Excellent Machinable (55%) (20%) (10%) (15%) Titanium Poor Excellent Stronger alloys Silver Good Poor Expensive Noble metals, Excellent Poor Expensive Gold Noble metal & (ie: Platinum-iridium 10%) Excellent High Expensive Gold alloys TABLE 3. -- ANATOMICAL SCREW COMPARISON 1. Cortical-ST 2. Cortical-NST 3. Cancellous 4. Malleolar Countersink 90= 90= Rounded Rounded angle Run out Short Short Long Long Point angle 60= -- -- 75=? Self tapping* Yes No No Yes *5-10% strength reduction axially. Note that cortical screws behave as lag screws only when they obtain a grip in the distal, and not proximal cortex. REFERENCES 1. Frankel, V.H. and Burstein, A.H.: Orthopaedic Biomechanics, Lea and Febiger, 1970. 2. Frost, H.M.: Orthopaedic Bimechanics, Charles C. Thomas, 1973. 3. Mears, D.C.: Materials and Orthopaedic Surgery, The Williams and Wilkins Company, 1979. 4. Swanson, S.A.V. and Freeman, M.A.R.: The Scientific Basis of Joint Replacement, John Wiley and Sons, 1977. 5. Vogler, H.W.: "Implants - Design and Nomenclature", lecture in Digital and Metatarsal Surgery, Pennsylvania College of Podiatric Medicine, September 25, 1984.