were accomplished by motion of 5° of flexion to 35° / of extension. Nearly all or these continuous tasks re~ quired only extension. Rising from a chair em- A ployed the gr~atest arc of motion, nearly 63°. Volz and coworkcrs (1980) also found that loss of wrist B mobility did not seriously impede performance of the activities of daily living, Volunteers with wrists Role of wrist position in finger function. A, Slight exten- immobilized in four different positions were asked sion of the wrist allows the flexor muscles to attain maxi~ mal functional length, permitting full flexion. B, Slight [0 rate their performance on 10 activities, and per~ flexion of the wrist places tension on the digital extensor formance averages were then computed for each tendons automatically opening the hand and aiding full position of immobilization. The results disclosed finger extension. the least compromise of hand function with wrists traL and 20 and 40\" of extension. Thev found that immobilized in 15° of extension (88% of normal grip strength was greatest at approximately 200 of performance) and the greatest disability \\vith the wrist extension and least at 40° of wrist nexion. wrists placed in 20° of ulnar deviation (71% of nor~ Vvith the wrist in 40° of extension and in the neutral mal function). position. grip strength was slightly less than the maximal values. Interaction of Wrist and Hand Motion. Studies by Hazelton and coworkers (1975) of the influence or wrist position on the force produced at \\,Vdst motion is essential for augmenting the fine the middle and distal phalanges revealed that the motor control of the fingers and hane!' Positioning greatest rorce was generated with the wrist in ulnm the wrist in the direction opposite that of the fingers deviation, the next greatest in extension, and the alters the functional length of the digital tendons so least in palmar flexion. Taken together, the results o that maximal finger movement can be attained. Volz and associates (1980) and Hazelton et al. \\Vrist extension is s)'nergistic to finger nexion and (1975) suggest that for grip to be effective and have increases the length of the finger flexor muscles, al- maximal force, the wrist must 1)c stable and must be lowing increased flexion with stretch (Fig. l4-22A) in slight extension and ulnar deviation. This conclu- (Tubiana, 1984). Conversely, some flexion oj\" the sion is consistent with the findings concerning load wrist puts tension on the long extensors, causing the transmission through the ulnar TFCC SLl1.1clures. fingers to open automatically and aiding hIll finger extension (Fig. 14-228). The position of the wrist also changes (he posi- tion of the lhunlb and fingers, thus affecting the The synergistic movements of the wrist extensors ability to grip. When the wrist is flexed with the and the more po\\vcrful digital flexors are facilitated hand relaxed, the pulp of the thumb reaches only by the architecture of the wrist. The digital l1exor the level of the DIP of the index finger; with the tendons cross the wrist within the depths of the wrist extcnded, the pulps of the thumb and index carpal arch and arc held close to the axis of wrist finger are passively in contact, creating an optimal flexion-extension, affecting wrist position l11ini~ situation for gripping or patching (Fig. 14~23). mally. By contrast, the extrinsic wrist flexors and ex- tensors moe positioned widely about the periphery to provide maximal morncnt arms 1'01' positioning the wrist. As the \\vrist changes its position ancl the func- lional lengths of the digital flexor tendons are al~ tered, the resultant forces in the fingers vary, afrect- ing the ability to grip. Volz and associates (1980) evaluated electrom~lographicallythe relationship of grip strength and wrist position. Grip strengths of 67, 134, 201, and 268 N were analyzed with the wrist in five positions: 40 and 20° of Ilexion, neu- '.\" ;'.
Man.\\· attempts h~l\\'L' hl..'L'!1 nwdc to Chlssif\\ dille orcnt patterns prL'ht,:nsik~ hand function. !'\\npi (1956) idclllified two distinct rmtlcrns or prchcn ;-;ilt...' lIlo\"elllent in thL' n(Jlnwl hand: pOWL\"r grip an pr~cbion grip. I-k (:mplwsizcd thai the fundalllen tal requisile 10 prL\"ilL\"nsioll, :-;Iabilit,\\', can bt,..\" mel b eilher posture. Power grip. or POWL'1' grasp, is a rorerI'll l act pc fnnncd with the finger lk'.'\\cd at all three joints s that the object is held bc!\\\\'een the finger and th palm, with the thumb positio!1ed on the palnmJ' sid or lile obje('l to fore\\! it sccllrL'I~' into tlw IXllm (Fi 14~14.·l). It is lIsliall.\\· performed wilh Ihe wrisl dev aiL'd ulnarl~-' and dorsillL'xed sljghll~' to <lugml'1H th lL'llsinll in the flexor (L'ndons. When the wrist is flexed, the tip of the thumb is level with the distal interphalangeal joint of the index finger. With the wrist in extension, the pulps of the thumb and index finger come passively into contact. Adapted with permission from Tubiana, R. (1984). Architecture and functions of the hand. In R. Tubiana, J. -M. Thomine, & E. Mackin (Eds). Exami- nalion oi the Hand and Upper Limb (.oP. 1-97). PhIladelphia: \\IV B, Saunders. Patterns of Prehensile Hand I. I\",-' Function ~ B Prehensile movements of the hand are those in which an object is seized and held partly or wholly I The two fundamental patterns of prehensile hand func- within the compass of the hand. Such movements are used in a broad range of purposive activities in- I tion. A, A typical power grip. The addllcted thumb forms volving handling of objects of all shapes and sizes. Efficient prehensile function depends on a multi- 1 clamp with the partly flexed fingers and the palm. The p tude of factors, the most important of which are: mar descent of metacarpals IV and V and additional flex~ ion in their respective Mep joints enable these fingers to l. Mobility of the first CMC joint and, to a lesser hold the object firmly against the palm, Counter·pressure extent, of the fourth and fifth MCP joints is applied by the thumb, which lies approximately in the plane of the palm. The wrist is deviated ulnurward and 2. Relative rigidity of the second and third CMC dorsiflexed slightly to increase the tension in the flexor joints tendons. Grip of an object along the oblique palmar axis (palmar groove), as shown here. involves a larger area of 3. Stability of the longitudinal linger and thumb contact. and thus more control, than does grip along the arches transverse palmar axis. 8, A typical precision maneuver. The object is pinched between the flexor aspects of the 4. Balanced synergism and antagonism between fingers and the thumb. The fingers are semiflexed and th the long extrinsic muscles and the hand in- thumb is abducted and opposed. The wrist is dorsiflexed. trinsic muscles Adapted wiriJ permission from Lands11lt'er. J.M.F. (1955j ,.'\\natomical iJnd {{mcrion,1} /fl\\festigalloflS Of) the articu/a!iDn D 5. Adequate senSOl)' input from all areas of the ;he human fingers Acta All,;, Suppl. 24. hand 6. The precise relationships among the length, mobility, and position of each digital ray.
Precision grip involves the manipulation of small A objects between the thumb and the flexor aspects of the fingers in a finely controlled manner (Fig. The \"dynamic tripod,\" a type of precision handling w 14-24B). The \\vrist position varies so as to increase the thumb, index finger, and middle finger work in c the manipulative range. The fingers arc generally in synergy for precision handling of the object while the a semiflexed position, and the thumb is palmady finger and the little finger provide support and static abducted and opposed. Certain prehensile activities trol. This functional configuration is demonstrated in involve both power and precision grips (Fig. 14-25). use of scissors (A) and a pencil (B). As a refinement of Napier's classification, Landsmeer (I962) suggested that the precision grip be termed \"precision handling\" because it involves no forceful gripping of the object and is a dynamic process without a static phase. In both power grip and precision handling, full opposition of the thumb to the ring and little fingers is obtained via palmar displacement of the metacarpals of these digits. A A variant of precision handling is the often \"dynamic tripod\" (Capenec 1956), wherein B _ thumb, index finger, and middle finger have namic action, \\vorking in close synergy for prec ,1m.IEI'-- handling of the object, while the ring and littl gers are used largely for support and static co The two fundamental patterns of hand function are used (Fig. 14-26). A further refinement is pinchi in unscrewing the lid of a tightly closed jar. A, As the mo- small object between the thumb and index f tion is begun, the right hand assumes a power grip pos- Such maneuvers are commonl).' classified a ture. B, As the lid loosens, the hand assumes a precision pinch, palmar pinch, lateral (or key) pinch, posture to perform the final stages of unscrewing. Adapted pulp (or ulnar) pinch, depending on the parts o with permission from Napier, fR. (7956). The prehensile move- phalanges brought to bear on the object being ments of the human hand. J Bone Joint Surg, 388, 902-913. dled (Fig. 14-27). Another important distinction between p grip and precision handling is the fundamen different position of the thumb in each postm the power grip, the thumb is adducted; in prec handling, it is pal marly abducted (Fig. 14-24) relationship of the hand to the forearm also d strikingly. In the power grip (Fig. 14-2411), the is usually deviated ulnarly and the wrist is hel proximately in a neutral position so that the axis of the thumb coincides with that of the arm. In this wa)', pronation and supination ca transmitted from the forearm to th~ object. In cision handling (Fig. I4-24B), the hand is gene held midway between wrist radial and ulnar d tion, and the wrist is markedly reflected in the
ture of the thumb. When the demand for prectSion A is minimuj or abscI1L. the thumb is wrapped over the dorsum of the middle phalanges of the digits and acts purely as a reinforcing mechanism. \\Vhen an el· cment of precision is required in what is predomi- nately a power grip, slich as the fencing gl\"ip (Fig. 14-28;\\), the thumb is adducted and aligned with the long axis of the cylinder so that, by means of small A Tip pinch i, B Palmar pinch Bc -, The fencing grip (A) is a power grip in which the elem of precision plays c1 large part. Instead of being wrapp _____;>.-~i/';' over the dorsum of the digits, the thumb is aligned w the long axis of the cylinder so that it can control the c Lateral (key) pinch tion in which the force is applied. In so doing, the thu loses its effect as a powerful buttress on the radial sid o Pulp (ulnar) pinch the hand. Hence, some of the power grip is sacrificed interest of precision. The coal-hammer grip (8) and th bunched fist (C) are examples of a string power grip w no element of precision. The ulnar deviation characte of a power grip is apparent in both cases. Adapted wit mission from Napier, J.R. (956). The prehensife movement the human hand. J Bone JOint Surg, 38B, 902-9/3. adjustrnents of posture, it can control 1hl: dire in which the force is being applied. Al the oth treme of the power grip range is the coal-ham grip (Fig. 14-28B), the crudest form of prehe function, where the thumb is wholly occupied orinforcing the clamping action the digits. A ample of this exlreme in an empty hand i bunched fist (Fig. 14-28Cl. -.~ Rotating the thumb into an opposing positi a requirement of almost every hand func Examples of precision handling in which small objects are whether it be a strong grip or a delicate prec pinched between the thumb and index finger. These grips pinch. In some inslances, howcver, the thumb not be involvcd at all, as in the hook grip in w are classified according to the parts of the phalanges the fingers are ncxcd so that their pads lie pa brought to bear on the object handled. A. Tip-ta-tip pinch. 8. Palmar pinch. C. lateral (key) pinch. D. Pulp (ulnar) pinch. and slighlly away from the p,t1m, together form hook. This posture requires relatively little m activity to maintain and is used when precisio
quirements arc minimal and when power must be gives the hand its great importance as an orga exerted continuously for long periods. Functionally, information and accomplishment. the hook grip pattern has limited potential and is not used ver~v often. An example of its use is to carry 10 The trapezoid. capitate, and second and an aoache case or a suitcase b.y its handle. In con- metacarpals, with their tight-fitting articulat trast, the individual whose hand intrinsic muscles form the immobile unit of the hancl. In their a are paralyzed or severely weakened relics on the ulation with the hamate, the fourth and hook grasp 1'01' all functional task completion. The metacarpals arc permitted a modest amoun hook grasp is the only grasp pattern available when palmar displacement, a motion essential for the hand intrinsics arc not working. ping. SUr/llnal'V 11 The most important motion of the thumb i position, in which abduction coupled with rot 1 The wrist is a complicated joint complex con- at the CMC joint moves the thumb toward the l sisting of the multiple articulations of the eight the lillie finget: carpal bones with the distal radius, the structures of the TFCC, the metacarpals, and each other. The 12 The finger rays are controlled by the cO carpal bones arc conventionally divided into a prox- nated action of the extrinsic and intrinsic mu imal and a distal row. systems. The operation of each ray is not compl independent of its neighbor's. 2 N1atians at the wrist include Ilexion-cxlension and radial-ulnar deviation. Stabilily during radial- 13 The components of the digital extensor as ulnar deviation is provided by a double-V system bly, especially the oblique retinacular ligaments formed by the palmar intrinsic ligament and the ra- count for the rdease of the distal phalanx and diolunate and ulnolunate ligaments. coupling of PIP and DIP joint motion. 3 The proximal and distal carpal ro\\vs form a 14 A unique feature of the MCP joints is bimuscular. bianicular chain that is subject to asymmetry, reflected in the bony configuratio collapse under compression. Stabilitv is provided the 111etacarpal heads, in the attachments of the by precise opposition of the articular surfaces lateral ligaments, and in the arrangement of th and intricate inlrinsic and extrinsic ligament con- terossei. straints. 15 The MCP joints arc stabilized primarily b 4 The extensor carpi ulnaris. e.'\\tensor pollicis radial and ulnar collaleralligamenl$ and also b brevis, and abductor pollicis longus act as a dy- transverse intermetacarpal ligament, which namic collateral system to provide wrist stability the palmar plates lo each other. during functional hand movements. 16 The digital flexor tendon sheath pulley sy 5- vVrist position affects the ability of the fingers is essential for maintaining a relatively con to flex and extend maximally and to grasp effec- moment arm for the finger Oexors and for mini tively. ing stress raisers between tendon and sheath. second and fourlh ul1tlldar pulleys playa par ,6 The TFee plays a significant role in cushioning larly imporLant role in this respect. compressive loads across the wrist joint. 17 The flexor superfkialis tendon has a gr overall excursion than does the flexor profun 7- The flexor carpi ulnaris is the most powerful The excursion of the Oexor; is larger than that o wrist motor and lends to place the \\vrist in a posi- extensors, and the excursion of the extrinsic m tion of flexion and ulnar deviation. tendons is generally greater than that of the in sic tendons. 8 The finger rays or the hand arc arranged in three arches: onc longitudinal and two transverse. 18 Extra excursion required at anyone joint Derangement or collapse of the arch system as a re- ing to disruption of the pulley system resuhs i sult of bone injury, rheumatic disease, or paralysis adequale excursion and subsequent weakness i of the intrinsic muscles of the hand can contribute more distal joints. to severe disability and deformilY. 19 The strength of the finger nexors is over 9 The hand is the principal instrument of touch. that of the eXlensors. The cOll1binalion of scnsibilit.y and motor function 20 Efficient prehensile I~mction depends on mobility of the thumb CMC joint and the fourth
fifth Me? joints, relative rigidity of the second and K.IlIl'r. LVI.G. (1980). Functional an.llom.\" of thl' wrist third CMe joints, balanced synergism-antelgonisl1l between the extrinsic and intrinsic muscles. and ad- Orrllop. 1-+9.9. equate sensDIY input. The relative lengths of the metac31-pals and phalanges and the (-Inger rays as a Kauer. J ..\\I.G. & Landsllll·er. J.,\\I.F. (1%1). Func whole are also important. ~\\Il:ll(.m~· of the wrbl. In R. Tuhinna (Ed.). Tilt· /lalld ;~tThc position of the thumb and the relationship between hand and forearm arc the most important I). Philadl'!phia: W.B. Saunders. differences between power grip and precision han~ eIling. Ku<:ynski. K. (1968). Th(.: UppC:f limh. In R. Passmore & Rohson (Eds.), ..1 COllllIW/;(J/1 10 Jlctlical SlIufic's (V Oxford: Blackwell Scil.:lltifi<.: Puhlications. L:llldsllll'l'r, L\\I.F. (1955). 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Applied Biomechanics ~ -,~ l:< \"'''~~'. -)00..
Introduction to the Biomechanics of Fracture Fixation Frederick 1. Kummer Introduction Fracture Stability and Healing Fracture Healing Surgical Factors Fixation Devices and Methods load Summary Suggested Reading .~'.
Introduction FRACTURE HEALING The study of fracture fixation biomechanics can be Currently, controversy exists about \\vhether com divided into two main areas: (l) criteria for achiev- pletely rigid fixation is the optimal condition fo ing fracture stability and promotion of bone healing bone healing. Ivlicromotion has been shown to a and (2) the characterization of techniques and de- healing. Healing results even in cases of gross mo vices intended to mechanically stabilize a fracture. tion such as that seen in rib fractures. Rigid fixatio An understanding of the biomechanical principles may lead to delayed healing, bone atrophy, and involved in these areas will aid the engineer in im- lack of external stimuli necessary for the healin plant design and enable the surgeon in selection of process. the most effective technique and device to obtain successful results in patients. Although gross motion between two or mo bone fragments usually leads to nonunion and Fracture Stability and brocartilage tissue formation, there is a low level Healing displacement (microlllotion) that appears advant geous to healing by providing a mechanical sign The clinical goal of effective fracture treatment is that stimulates the biological repair processes. Th rapid healing, withollt significant deformity or li1l1b amount of local strain in the healing region (chang shortening, to restore the patient to a pre-fracture in length divided by the original length) seems to d level of function. In the elderly. rapid mobilization is termine the nature of the tissues formed (e.g., fibr essential to prevent deleterious consequences of bed cartilage or bone). The optimal frequency, \\vav stay. The first goal of treatlnent is fracture stabiliza- form, and total number of cycles of this signal cu tion. This is determined by the location and type of rently are being investigated. Several methods fnlcture, the muscle and body forces acting on it, and promote healing by externally stimulating a fTactu the various passive soft tissue constraints such as lig- with ultrasound or electromagnetic fields are aments and fascia. Some simple fractures are inher- clinical use (Case Study 15-1). Concern also exis ently stable with low loading and thus require nlini- regarding the process of stress shielding that occu mal treatment, such as sling (clavicle), while others, when the fixation device carries most or all of th such as midshaft comminuted femur fracture, require mechanical load and thus, by Wolffs law, promote major surgical intervention and insertion of an inter- localized osseous resorption as a result of the resu nal fixation device for adequate fixation. Although an tant unloading of the bone around the device. Th osteotomy (a surgically created fracture for correc- is often referred to as load bearing versus load sha tion of defonnity) allows close approximation of frac- ing. Much of the initial osteopenia seen benea ture ends, typical fractures are often fragmented and fracture plates, howevel~ is thought to be caused b usually lack inherent stability. The interdigitation of vascular disruption during their application. the bone ends can facilitate stability, such as when a tapered bone end is inserted into the nledullar:v cavity! Bone healing in the presence of a gap with lTIin (creating a displacement deformity, however). mal movement passes through several stages of r pair \\vith a concomitant increase in mechanic Traditional methods for the treatment of frac- strength as mineralization increases: hematom tures are externally applied and include traction, and inflammation, callus formation, replacement b cast, and braces. External forces or constraints ap- woven bone, and fmally remodeling into lamellar plied to the injured limb act to stabilize the fracture trabecular bone. Callus can~form both periosteal (by limiting ITIuscle or soft tissue forces leading to and endosteally and enlarges the diameter of th deformity) and maintain alignnlent of the limb. bone at the fracture site. Although callus is le Howevel~ in many cases, as a result of the nature of strong and stiff than nlature bone, this increased d the fracture or the patient's condition, an internal or ameter can increase stiffness in bending and torsio external fixation device attached directly to the bone at the fracture site as a result of the increased m is required to achieve adequate fracture stabiliza- ments of inertia. Direct bone apposition by com tion. The designs and use of these fixation devices pression with rigid fixation, in which the initial r rely on an understanding of bone healing and the pair stages are eliminated or minimized, heals by loads and forces to which the device is subjected. remodeling process and can take longer becau vascularity must be re-established. The other impOl·tant factor for healing is ad quate blood supply that necessitates the surgeo 39
Ultrasound Treatment for Fractu\"e Healing acoustIC rodiaHon at irequ€-nCles above the !lm!! of hu- man heanng. Its acoustIC raeJlCIllon, In the form of pres- A40-year-old female involved in il mOlOr vehicle collision sure waves, provides rnICfOrnf.'chiH·\\ICa! stress and force' to the bone ,me! surrounding tiSSUe>, ThiS rnedh1nI((ll stunu· § _ in December sustained a left tibiofihular lr<1Cwre treated lalion plays a major role in bone hedlll1g because bone: with external fixation. Case Study Fig. 15-1·1, In January, lo'.>\\'- reacts to the amount and dir(;ction of force and remof.!· intensity pulsed ultrasound (US) was mitiatecJ 10 promote frac- els to adapt to the applied suess \"mel its elirf~c!i()n ture healing. Case Study Fig. 15-1-2, in March, 3 months postfracture and 2 months after mitiation of pulsed US appli- ReprHlted V-11th permission from v\\lolH, J (1986). Das cation, early healing is detected (arrow). Case Study Fig Geseu der Transformarion <ft.'r Knochen [The law oj 15-1-3, in May, 5 months after injury and 1\\ months follo\\,..,mg bone remodeling]. P. Milquet & R. Fulong (Trans.). Berlin: initiation of US, the bone healing is successful. Springer-Verlag. (Oflglllal ..vork published in 1892). Pulsed low-intensity ultrasound has been succEssfully used for fracture repair (Frankel. 1998). Ultrasound is an presel·ving the vascular supply' of the bone (e.g., pe- b~nding! and/or torsion) and magnitude of force riosteum) and providing conditions for early rcvas- which the fixation will he 'subjected and whe cularization by; careful opcrative techniquc (e.g.! these forces will be cyclic, requiring ~\\dditi sort tissue presenTation). Numerous studies 1lave strength of fixation 10 ~\\ccnunt for possible dC\\'ic demonstrated a direct relationship between the tigue (Case S(ud~' 15-2). The second factor is quantity and quality of microvascular structures in bone quality', which delermines the strength a\\ the healing region and the rate of ronn~llion and re- sultant mechanical properties of\" lhe new bone. able to supporl the fi:..;atioll dc\\·jcc. Other factors related to surgical and ~lIlatomical considerati SURGICAL FACTORS for example, the expOSllre (possible scarring and cular compromisc), wIlL,thcl- the devicc will fit Many factors determine the optimal fixation method for a specific fracture appliGltion. First arc mechan- quately within the soft tissuC's. and if nL'urovasc ical considerations, specif-lcally the types (tension, structures are at risk. The nature of the origina .iur:v and the amount of soft tissue d~IIll~lgC also tern1illC treatment l~chl1iques. )1 '; t\"h.,_ '. ,
lillIiVIEf.1C-------·--··---·----- ··-·--r (number of patients, adequate follow-up) so that th many vadables can be properly analyzed and Fixation Plate Failure proper statistical significance determined. .' n internal contemporary fixation plate inserted into Fixation Devices A the arm of a 25-year·old male who sustained a frac- and iVlethods ture of the radius. The plate was fractured as a result of vVires, staples, pins, plates, and screws are the com fatigue 20 years later. Repeated loading and unloading of mon implant devices used to achieve fracture fixa tion. These are usually made of stainless stee a material will cause it to fail, even if the loads are below (316L), sometimes of titanium alloy (Ti-6A l-4V), o occasionally of cobalt-chromium alloy'. Each meta -.: .the ultimate stress (Simon. 1994). Each loading cycle pro- has advantages and disadvantages such as strength duces a minute amount of microdamage that accumulates modulus (stiffness), corrosion resistance, and cas with repetitive loads until the material fails. Mechanical of imaging (MRf, CAT). Sometimes there is a \"race considerations as to the magnitude and repetition of the between healing of the bone and fraclure, usually b loads to which the fixation will be subjected should be faligue, of the device. There is a recent interest i considered, along-with the fatigue life of the material. This the clinical application of biodegradable polymer is recorded on a curve of stress versus number of cycles. such as polylactic acid. Polymers are more Ocxibl T,hus, higher stresses produce failure in fewer cycles (load· than rnetals and would lead to greater load be:1l-in ing to the ultimate stress produces failure in one cycle>. by the healing Ii·acture; biodegradable materials d while lower stresses are tolerated for an extended period. not have to be removed in a secondary' operatio and their mechanical properties gradually decreas Case Study Figure 15-2-1. with time, thus avoiding stress shielding. Ho\\Vevcl their mechanical strength is much less than that o Evaluation of fixation strength can be accolll- metal and some of the degradation products hav plished by laboratory testing of actual implants in shown untoward biological responses. Researc cadaver bone. One difficulLy of such testing is in ad- continues into the use of various glues, cements equately simulating in the test model the complex in and adhesives for fracture fixation, some of whic vivo, cyclic forces on the device. Another difficulty is are also biodegl-adable. in simulating the biological repair processes that would act to stabilize the fixation over time. Ca- Wire fixation lIsed as cerclage or a bone suture i daver studies also can determine the anatomical a common application; in both cases multiple wire structures at risk. Computer modeling such as finite are required to provide stable, three-dimensiona element analysis can be lIsed as an initial method to fixation. This requires achieving equal tension dur evaluate fixation methods and device designs but re- ing tightening because loosening at one or mor quires quantified parameters of bone modulus and sites can provide a locus for motion and possibl strength, which may be lacking. for an exact solu- nonunion or cause malpositioning. Problems wit v.lire fracture fixation include the necessity and sur tion. gical complexity of making a'bone hole and passin the wire, breakage during tightening or later as a rc Clinical trials is the other major method used to suit of fatigue (cyclic loading), and cut-through o evaluate the efficacy of a particular fixation method. the bone. For cerclage applications, there is concer HOWe\\lel~ care must be taken to adopt the appropri- about compromise of the periosteal blood suppl ate techniques to quantify data and design the trial and the resulting increased healing time require for revascularization. Some~recent developments are wire tensioning twisting instruments and the usc of crimping sys tems to avoid the problems with twisting or knot ty ing. There are also new oriented polymers (Spectra that do not stretch as traditional suture malcrials d that can be Llsed with a suture anchor system t
eliminate the difficult.\\' of looping a suture through /------ ...... bone and suture abrasion against the bone. \\/ Staples alone usually' do not provide sufficient mechanical stability for permanent fixation and ~\\ / I their use often requires pre-drilling holes for the sta~ \\\\ I pic legs; pneumatic-driven staples can be used to rapidly tack fragments prior to a more rigid fixation \\ \\ /1 but need careful control of the insertion driving force to prevent untoward damage to the bone. \\ \\ /1 Some staple designs have been developed that can effect compression during insertion, such as staple \\ '1/ fabrication from nitinol (an alloy.' that changes y~ shape when heated to body temperature). / I--il- Tension b Kirschner \\vires are normally used to hold frag- nlents of bone prior to rigid fixation and for percu~ /1 wire loop taneous pinning of small bone fractures but, in gen~ eral, lack sufficient mechanical stability for their /I use as a primary fixation in weight-bearing bones. At least two wires should be Llsed for each bone // fragment, and they should not be inserted in a par- allel manner to prevent \"pis toning\" of the bone frag- / ment along the wires (Fig. 15-1). Threaded pins pro- vide additional stability because they minimize Tension band wiring of two K-wires; tightening the wi sliding of the bone fragments, but their rcmoval is loop applies compression to the fixation. K-wires are in morc difficult. Occasionall'y', pinning is uscd in com- serted in a skewed configuration for stability, bination with sutures looped and tightencd around the pin ends or through loops in the pins. This \"len- portant to prevent stripping of the bone or sc sion band\" techniquc provides significantly' in- head torsional Failure. creased 111echanical stability of the fixation. Because of anatomical constraints or surgical The major intrinsic factors that influence scrc\\v- posure, screws cannot <.lIways be inserted perp holding po\\ver are outer thrcad diameter, thread dicular to the bone axis or the orientation of configuration, and thread length; extrinsic factors ends of thc bone fragments is not perpendicula are bone quality, typc, and screw insertion orienta- the screw axis. In this case, the holding power of tion and driving torque (Fig. 15-2). The two basic screw is decreased and a shear component of types of screws are cortical and cancellous and are holding Force is created, which acts to destabi distinguished by their thrcad design; cancellous alignment. Pre-tapping of thc scrcws usually! is screws have a greater distance bct\\veen adjacent necessary and has been shown to have minima threads (pitch) and the ratio of outer thread diame- fect on their holding ability; many' screws are ter to body diameter (Fig. 15-3). tapping owing to a modification to the design of leading threads. Usually lwo or more screws are The inherent holding power of a screw is a func- quired for functioning, although one screw has b tion of the outer thread dianlcter times the length of suggested for some applications if sufficient in its threads within the bone. When used to hold two bone fragments together, scre\\vs comnlonl:v are used in a lag modality in which the proximal por- tion of the screw remains free within one fragment (either b:v using a screw design having no proximal threads or b)' enlargement of the hole in the proxi- mal fragment, which should require thc use of a washcr under the screw head for adequate support). Insertion torque detcrmines the force with which bone fragments are held togethel~ which, in turn, crcates the friction that prevents their motion. Con- trol of torque (torque-limiting scre\\vdrivcr) is im-
Outer thread ~I rragment approximation can be achieved to create diameler adequate friction between the bone surfaces for sta- bility. The quality of bone also determines serew- Length Thread holding ability; cortical bone is approximately ten of screw pitch times stronger than cancellous bonc. The thickness in bone' of the cortex and the degree of osteopenia (bone density) are thus critical for fixation strength and in- ,. nucnce the number of screws required for adequate stability. Using screws in a biconical manner appre- Screw parameter~. For ~crew pull-out the bone must shear ciably increases the strength of fixation. along the outer diameter (dotted line). Anatomical constraints limit the number or size of screws that can be applied in a given region. As a result, screws are ohen combined with plates to achieve adequate stability and increased strength of fixation. The optimal site for a single plate applica- tion is on the side of the bone subjected to tension; lIsually two plates are applied to achieve better FIxa- Lion stability (Fig. 15-4). Owing to anatomical COI1- strail1ls such as soh tissue thickness. sometimes thinner plates are used (such as for forearm fracture stabilization), but these plates possess sufficient stiffness (function of the width of the plate times its thickness cubed) Lo prevent undesired fracture mo- tion as a result of bending loads (Fig. 15-5). Screws should be inserted with a torque driver and the tightness of all screws rechecked; if this is not done, one screw may bear most of the load and possibly Fail. Some plates use a specially designed scre\\v hole slot, countersunk Lo accommodate the Load Load ~~ Plate Fracture Gap_ Plate gap closes ope~s -~'!lIlIJ'llm~ Plale on Plate on tension side compression side Types of bone screws, left to right: cortical, cancellous, Effect of plate placement. A plate located on the compres- and cancellous lag. sion side causes the fracture to gap when loaded.
Posterior Posterior thl.'ir insertion and the possibilit.\\\" (II\" compromisin bending force p\\.'riosh.';.11 blood stlppl.,· b.\\· thL' l':\\posure or pl;'11(' i sert iOl1 (somi..' plati..' designs have inferior fed O I ridges (0 minilnize this possibilil:!). Therc is ;.tlso i lCrest ill pol:·tTlcric plales lhat would he tTlOrL~ flex Lateral ble 10 achicvc a grealt:r degre\\.' of Inicromoti(l!l till.\" fracture:.: thai could he:.: advantageous 10 bon 1.8 em Medial healing and minimiz\\.' stress shielding. - bending force Hip fr.:Ictlire ck\\'ices call be intt.·rnal orexlcl\"nal thl.'ir application; lilt: most cornlTlon \\.'xternal dt.:\\'i Anterior is a side.: plate affixed to the felllur supponing an i Icrn[ll. sliding hlg screw through the fleck across th Effect of loading direction on plate stiffness. The rigidity of the plate is EI, where E is the modulus of the plate mao fmctlll\"c. terial and I is the moment of inertia of the plate. I == bh1/12 ( II = posterior bending; 11 = medial bending; II = The relationship uf biomt'chanical forces born h,,· the tk'\\'ice ;'Hld borne b-,· the bone (load \\.x'arin 0.5 x 1.8)/12 = 0.243; I~ = 1.8 x 0.5J/12 = 0.01875), where b load sharing) inllllt.'nces fracture hc;;ding ;'\\lld devi sun.'h·al. An important ractor is the.: ability 01\" the d is the base dimension and h is its height. Thus, the plate is vice (0 slide 10 cOllsolidate the rnlClurc during hea 13 times more rigid in posterior bending than in medial ing. lnt\\..'rnal c!t:.'\\\"ic(':-; for fix~ltioll arc llsually i bending. traIl1L'dullal\".\\· nails (Ii\\,1) alld in cOinparison wi ('x (('1\"1141 I dc\\\"iccs arL' SUhjL'ct to less loading fore\\.' b screw head, whose center is offset with respect to the screw head to obtain interfragmentary compres- C;;luse or their location being closer to t1h.. Ileutr sion as the scrc\\v is tightened. An alternative method to achieve compression is to pre-bend the bending axis 01\" the bon\\..' (Fig. 15·6). Their size plate before application so that when the attach- ment screws are tightened, the bone fragments are critical bcc..wsc their ht..'nding and tor:-;ional stifTnL' approximated as the plate straightens. Some new ;;lre proportional to the diameter to tile four plate designs use threaded holes to engage the screw power. This is \\\\'11.\\· (me large nail provides trw so that bicOl-tical screw insertion is not as essential rigid (jxalion than docs lllllltipk smaller rods. Siz for maximum fixation stabilily. or<:lmoullt ClilTaturc, and amount 01\" n:~;'lIning a Plates can also be used to span gaps created by severe fractures or tumor surgel), and are frequently also important because the swbilit.\\\" of the fixatio used with bone grafts for this application. Unless the graft is exactly sized. the plate will bear the en- rL'lics on lond transfer to the bont.~: often, distal an tire load across the defecl. The bending moment on the plate-screw fixation linearly increases with c1e~ Typical intermeduliary and extramedullary devices. Top, fect size and thus the plate requires adequate stabi- Medoff sliding plate; bottom, intermedullary hip screw\" lization, particularly at the morc highly loaded, proximal end where at least three screws are • needed. Multiple-holed plates enable selection of the best osseous sites 1'01\" screw purchase and should permit anchoring of the graft by at least two addi- tional screws. The major surgical considerations I'm\" the lise of plates are the requirements of a large exposure for
proximal screws inserted through the bone and nail Type olframe and arc used to increase torsional stability (Fig. 15-7). connections Bending of the nail as a result of insertion in a curvcd medullary cavity can make inserlion of the Size and distal screws difficult. pin type External fixation devices are also used for fnK- 11 ture stabilization; multiple transcutaneous pins are inserted into the bone and stabilized with an exter- 1 Pin I Fracture I nal bares) or ring(s). Factors that influence mechan- distance spacing ical stability and rigidity of these constructs are the numbel~ diameter, orientation, and length of these Typical external fixator showing the variable that influ- pins and their relation with respect to the fracture: ences fixation stability. however, these factors arc subject to surgical con- siderations for frame application. Large, short pins tion of appropriate sites for device attachmenl to located close to the fracture provide the most rigid th~ spine are IitrJi ted. This is imponanl because fixation (Fig. 15-8). lhese devices are subject to appreciable forces dur- ing flexion and extension of the neck, and torso and Spinal implants used for deforrnity correction or fixation failure can occur. fracture fixation consist of various combinations of rods, wires, plates, and screws. The junctions be- tween these components are often the site of failure, such as fatigue or fretting as a result of cyclic load- ing. A specific problem is that the size and the loc3- Summarv Greater 1 The optimal frequency of waveform and the to- bending tal numbcl· of cycles of bone healing are under in- vestigation. Rigid and semirigid (micromotion) de- \" Higher vices can result in bone healing. Gross motion, compressive instability, and inadequate blood supplv may lead to loads nonunion. Extramedullary Intermedullary 2 Factors determining optimal fixalions for spe- cific fraclure applications are: !;... a. Mechanical considerations, such as the The extramedullary device is less rigid and when loaded types and magnitude of forces to which the fixation will be subjected and expected has greater deflection, creating higher medial stresses in cycle histOl\",!1 the femur. b. Bone qualit'y and st,:ength c. Surgical and anatomical considerations d. The energy involved in the original injury and the amount of soft Lissue damage. SUGGESTED READING Burstein, A.H. & Wright, T.M. (1994). FItIll[all/i:J1/a[s of 0,.· tltoflW.'dic Biol1lcc!ulIlic:s. Baltimore: Williams & Wilkins. Codlran. G.V.B. (1982).:1 Primer oJ' Orthoplledic lJiOIllCc!UlII- it..'s. London: Churchill Li\\·ingstonc. DYi.:hc\"IlI..'. P.. &. Hastings. G.W.-(Eds.) (l9S .l). FllIlc:tiOlud Be- . hm;;or of Orlhop(ll'dl~· !J;omata;als. -Bot-'I Raton. FL: eRe Press. ~\"\"'===~====~===\"\"'==~=---=\"\"\"~\".~\",.\".\"_\"~\",_\"\"'~\"--======\"\"',\"\",,;r,- '0'J!'G'7,;- ;\"c-~\"'<,!\",:\"''_..,-,'\":_'r.\".··\"\"\",,.....\"......,,:::s;-\"'H-\"~,_
Fr~lnkd, V.H. {I 99S). Resldts or prescription or pulse ultra- Radin. E.L. (Ed.) (1992). \"raniral BioJllecJlllllics (01' the 0 sound therapy in rl'.u.:tun: management. In Z. Szabo, J. tJwfll.'liic Surgeol/. Nl:\\\\' York: Churchill Livingstone. L~\\\\'is, G. F311tini. 0: R. Sa\\·~tlgi (Eds.), 5111\"gi('(// 7t:dlllology Sarmiento. A., McKdlop. '·I.A., Llinas. A., c[ al. (1996). Eff IUlc!t\"/WliOlw{ VII. IJll~n/(/tiollal Dct'dlJIJllli.!lIls \"11 SlIrg{.'(\\· of IO<:lding Hnd fraclUn: motions on diaphyscaltibi.l! fr tureS. J OrtllOp Ues, /.+( I). 80-84. (//ul Surgical R,:s{.'(/J\"l.'h. Snn Francisco: Universnl ~lcdical Press. Inc. Surgical Technology Inlcrnationnl. Simon. S.R. (1994). OrrJwpcdic Basic S(.'iclln:. Rosemolll. ,I,IOS. Fun~, Y.c. (1981). BiomecJulllics: ,\\-lcdulI/lt:a/ Properties of Sph'ak, J.~'l., Oi Cesarl:, P.E .. Feldman, O.S\" ct til. (Ed L-it'illg TisSlil'S. Ncw York: Springcr-Vcl'bg. (1999). Orthopacdics-A Suu/y Gllidi.'. New York: McGra Hill. Ghista, D.N. (Ed.) (1981). BiolllecJUlllics o{ Medical Dn'ices. Tencer, A.F. & Johnson, K.D. (1994). 13iolllecllllllics ill Or/h Nt\\\\' )'ork: Dekkt:r. orfl,'dic 7hlllllUI. London: M. Dunitz.. orComa, E.R., Harrington, I.J., & Evans, D.C, (1982). Biollle- duwi\"s AJlIsclilo.d:dcUI{ 11I;llry. Baltimore: Williams & Uhthoff, H.K. (Ed.) (1980). C\"rrclI! COl/cep!s Jl1rcn/{l{ Fi liol/ 0\" Frac!ures. Berlin: Springcr-Vtrlag. Wilkins. Kasscl', J.ll.. (Ed.) (1995). OrlJwpai'dic Knowledge Update .i. Valcllw. J. (Ed.) ( 1993). 8iollle(.'I1(l1Iio. Amsterd,un: Elsevie White, A.A. &.: Panjabi, M.M. (£ds.) (1990). CliHical Biol RoselJl<)lll, IL: AA05. Mow. V.e. 0: l'layes, W.e. (Eds.) (1991). Basic Orthopaedic cJWI1It.·s o{ the Spiltc. Phil;:lddphia: Lippincott. Wolff, J. (1986). Dos Geset:. tier Trtms{o1'l11l1tioll del\" Knoch BiOllh'challics. Ncw York: Ravcn Prcss. [The 1;:1\\\\' of bone l\"enloddingJ. P. ,\\,Iaqul:t & R. Fulo Ozka\\'a, N. & Nordin, M. (1998). Fllluhmll'ltrals o{ Biome- (Trans.). Berlin: Springcr.Vcr!;g. (Original work publish d;aJlio (2nd .::d.). Nl'W York: Springer-Verlag. in 1892). Pauwels, F. (1976). BivJIlt:clulllics o{ the NOI'I//(/i ami Dist:(LWd fljp. Berlin: Springl,.·r.Vcriag.
Biomechanics of Arth roplasty Debra E. Hurwitz, Thomas P Andriacchi, Gunnar BJ Andersson Introduction Forces at the Hip Joint Rotational Moments About the Implant Reconstructed Joint Geometry Stem Position Within the Femoral Canal Periprosthetic Bone loss Forces at the Knee Joint Medial-lateral load Distribution Patellofemoral Joint and loads Joint line Height Posterior Crucfate Ligament Conformity Constraint Polyethylene Anterior Cruciate Ligament Summary References
Introduction To meet one design objective, compromises with r spect to another design objecLive must often b As designs for total joint replacements evolve, there made. For example, the femoral stem of ccmcntle is an increasing need to improve the understanding total hip replacements requires initial rotatory sta of the biomechanics of these joints. For instance, bility for bony ingrowth into the porous material t reducing loads on a total joint replacement during occur. The rotational stability of early ccmentle daily activities and designing implants to withstand total hip replacements was insufflcient and was in these loads will both increase the likelihood of nor- creased by designing implants with greater mcdu mal joint function and lower the probability of im- lary canal fill. Ho\\vever, implants with increase plant failure. As implants arc put into younger and medullary canal fill were stiffer because of the more active patienls, the IO~lds placed on the illl~ larger size and were thll:) associated with great plants arc often several times greater, thereb.y in- stress shielding. creasing the likelihood of mechanical failure. Me- chanical problems associated with lotal joint Forces acting at the hip or knee joint arc depen replacements include issues related to wear of the dent on the external forces aCLing on the limb an bearing surface, mechanical failure of the implant, the internal forces primarily generated by musc loosening of the implant from the bone, and dislo- contraction. Joint forces have been measured wit cation of the implant at the articulating surfaces. implanted transducers or estimated using invers Important biomechanical infOl-mation addressing dynamics and analytical methods (Tables 16-1 an eHch of these issues has been obtained from in vitro 16-2). By measuring the external ground reactio cadaver testing, model simulation of the joint mo~ forccs, approximating the limb segment inerti tions and loads, and in vivo studies of human loco- properties, and localing the three-dimensional pos motion. tion of the joint centers during dynamic activitie the intersegmenLal hip forces and external momen The overall goals of joint replacement are to pro- can be obtained using inverse dynamics. Even \\\\li vide long-term restoration of function and pain re- simplifications to the anatomy, solving For the mu lief. Several mechanical challenges must be met to cle and contact forces at the hip and knee remain achieve these goals. For example, the control of an indeterminate problem. Thus, a unique solutio joint motion and stability in total knee replace- for individual muscle forces is not possible withou ments is achieved through an inLl'icate biomechani- further assumptions or simplifications. cal interaction between the shapes 01\" the replaced In general, two approaches have been used solve the indeterminate problem. The first is a \"r surfaces and the remaining ligaments and muscles. Hip and Knee Contact Forces Measured In Vivo in Patients With Instrumented Implants Activity Typical Peak Force Number Time Since Reference (Body Weights) of Patients Surgery (Months) Kotzar et aI., 1991 Davy et a\\., 19B8 Hip Force 2.7-3.6 2 1-2 Bergmann et a\\., 1993, 199 Walking normal to fast speeds 2.6 Stair climbing 2.6 2 8-33 Rydell, 1966 Walking slow speed (crutches) 2.6 2 English & Kilvinglon, 1979 1 6 Taylor el al . 1997 Ascending stairs 2.7-4.3 3.4-5.5 t5 Walking 3.9-5.1 12 Ascending stairs 1.8-3.3 Descending stairs Walking 2.7 Walking slow speed Knee Contact Force 2.3-2.5 Walking normal to fast speeds 40 -',--; /'-~-
~------------ Analytical Methods of Estimating Peak Hip and Knee Contact Force Activity Magnitude (BW) Method Reference Hip Force 4.8 Reduction nlE.'thod P(1UI. 1976 Walking 7.2 Reduction tneihod Pc)tJI, 1976 Stair ascending 7. , SI'lIr ciescendlllg 2.2 0plimILdtlon-maXlfniZe Brand & Crol\"vnmshield, 198 VValking slow w!th cane 3.4 endurclnce VValkltlq 510\"\"./ wltllOlll cane Cro'lojl)H1Sflleld 0t aI., 1978 vValkinD 3.3 QuasI SIc1!IC-!llmm)lb2 muscle lorces Stair climbln~l 5.S Selreg & ArvJkar, i 975 (hal( rising Re<.i~:ctlon method Walkmg 32 Rt~uuctlon method Harnng!Oo, 1983 Op!imil<:tlon Knee Force 4 Optin11Z,ltl0n i'-Aofflson, 1970 VV,llklng 3 Sclllpp!eJn fit Andric1cchi, i99 Walking 7 5(>lr09 & ;\\rvlkar, i 975 Walking W,11kin 9 dUClion\" method I1wl groups llluscles into func- Forces {{t the Hip Joint tional units, tht:reb~' reducing th~ indcl\\.'rminall' problem 10 a dC(l..'l\"n1inak olle. The second ap- The pt.'ak resllll~lI1t forces during gait 11l('asurc.:d wi proach lIses optimization methods in which the strain-gauged pro:·ahcscs han: ranged 1'1'0111 1.8 force is distributed among lIw muscles in such a ..t.36 limes bod~' wdglllS (T~lbk 16-1 l. Bergmann lllanll\\.'I' thaI so III \\.' ph~-'sicnl paralllClL'r stich as al. (1993) havL' conducted tile most CXlL'llsivL' in vi minimum muscle stress, maximum muscle cn- studies on rO]T~S during dail,\" aCli\"ities (Fig. 16- c1ur::\\Ilcc, or minimum joint reaction force is opti- and have SIlO\\\\'1l tlwt tile p(~~lk i<)ads and torsion mized (Brand, Pedersen, &: Friederich. 1986; lllOIlH.:nts im.TL'USC with w~dking speed. In genera Crowninshield & Brand, 1981; Cro\\\\'ninshidd ci the force at IIlL' hip joint rL'~ldh..'d all initial Ik'ak al.. 1978: Scircg & Arvikar, 1975). How('\\'L'r, func- early stance and a second peak in lale SWIlCC. J\\tla tion in patienls with lOW I joinl rL'placcmL'n(s ma.\" or the failure nH.'d\\~\\Ilism~ in total joint rL'\"plac not be normal and ma.\" b(' innue]H.:ed less b~' miJl- llh.;nts an.: ark'rled nol only b~' load magnitll(IL' b imizing L'\"nergy than by olher fnclors such as pain <llsu by IIll' c~'c1ic nature oT till,.' 1()~l(ls. Thus, in L'\\' or implant slabilit~'. One altcrllatin~ 10 optimiza- unting c.n::lic fatigue, one must consider (wo cycl lion mcthods is 10 generate a range or muscle of load \\'ariat ion rrH\" each Sll'p. fortes able 10 maintain mechanical equilibriulTl at IndqK'ndent rnL'aSUrell1t.:.'nts at diiTen..'nt instit the joint. Rather than obtain a single value for lltc tions seClll to indicate tllat st~ms with silllilar featur contact or muscle ['orcT, this parametric approach l,.·xJ.x:rience similar peak forces. fn rnan,v illsl~nlCc delcnnincs a rnng(.· of agonisl muscle force distri- diffcrL'IH.:es between the IllL'asurcd forces can be a butions needed (0 balance tile L'x(ernal moment lribUlcd to variations in the hip position during ga and a selected level of antagonistic actl\\'Il~; (Nobk, l-Ie1mk~, & Paul, 1993). Anakti<:al mode (Fouch~r ~t aI., 1997; HUlwitz, D. E., 199~. Func- gL'l1t.'rall.\\· predict higher contact forces titan lho or,, tional Biomechanics tilL' Hip As RL'l~llcd to Total measured b.\\' in \\'in) instll.lIlH..'lHcd total hip replac Hip Replacements. UnpublishL'd d()ctoral dissL'rla· IllL~nts. This Illa~' IlL' IllL'\" rl'sult of employing anatom lion. Univcrsi(~1 of Illinois at Chkago). cal simplirlcations, using inndequHtc optimization <.
A Activities: Walking, Jogging. Stairs, Stumbling teria to determine the agonistic and antagonist muscle activity, and assuming norrnal function. Ide 900 RESULTANT FORCE R tifying how patients with total hip replacements alt their function during gait could irnprove analytic 800 EBl EBR JB predictions of joint force. In vivo force measuremen have often been obtained during the early postoper 700 tive period \\vhen patients may still have compromise function. ~ 600 'u, 500 Rotational Moments About <D the hnplant t 400 I /. AI Out-or-plane loads l11ay be detrimental to both in a: 300 I II II tial and long-term implant stability, especially ullcemented stems (Berzins, Sumner, & Andria 200 chi, 1993; Phillips, Nguyen, & Munro, 1991). E cessive bone-implant motion prevents bone i 100 ESJ 6Mmli<ln Abs. max growth into porous coatings and may lead failure of the biological fixation (Pilliar, Lee, 0 ,~ \"~ % 0 ~ ~~ ~ %~ ,0 Maniatopoulus, 1986). The largest torsional m ments measured in vivo during activities of dai ~ .\"§ 6 3 'i 'i , .§ 0 \"~ living (Bergmann, Graichen, & Rohlmann, t 99 \"2 w reach the average experimental strcngth of impla w § M \" ~' \"\" M2 fixation (33.1 Nrn) as dctermined from in vit tests (Phillips, Nguyen, & Munro, 1991) (Fig. 16-1 ~ ~ \" \" \"~~ E \"~.< ~ In vitro testing of prostheses implanted in huma w femurs indicate that the amount of initial bone-im plant motion is sensitive to the off-axis loadin B BENDING MOMENT Mf that frequently! occurs during stair climbing an rising from a chair (Fig. 16-2) (Berzins, Sumnec I 18 1\\ Andriacchi, 1993). Furthermore, resultant prox 16 mal translations were twice as high with a straig :i: 14 1\\ 6 D stcm as with a curved stem at load angles encou 12 tered during these activities (Berzins, Surnner, <D 10 A~ A Andriacchi, 1993). 8 .0 6 d r 61 \\Valking with a decreased hip range of Illoti 4 during daily activities may minimize the out-o o· 2 Pli )11 plane force components (anterior-posterior). Thu decreased sagittal plane motion during walking o ;; a ten reported in patients with total hip replac ments (Hurwitz, Chertack, & Andriacchi, 199 ESJ MiJd'i1rl D. Abs. 01i1~ Murray, BreweI~ & Zuege, 1972; Stauffer, Smidt, Wadsworth, 1974; White, Yack, & Lesswing, 199 f ,~0Fc ;0:; ~ •\" •~ \" , ~~€ g. 0 % may be beneficial for iIllplant stability b,Y reduci the rotational moments about the implant stem. D 0 :i 11 namic hiii range of motion during gait in patien with uncemented total hip replacements has be , \" \"'0 0 M §\" related to the metaphyseal fill of the femoral ste (Hurwitz et aI\" 1992). It was concluded that a r ~~~ ~ ~ M ~~ ~w duction in the dynamic range of hip motion \\vas ~~ ~ adaptive response to decrease torsional microm w5 ~ \"~ if\", ~ ~.0 c TORSIONAL MOMENT Ml 7 6 6 I5 A A AA A :i: 4 6 u, 6 e:<D I F, -- 3 II ;; 2 o Meclian D. Abs, 01i1~ !n -- o -- Joint loading during different activities for two patients, EB (EBL: left hip and EBR: right hip) and JB. The median load val- ues (column height) and absolute maxima (triangle) are given. The resultant force is in the top graph (A) the bending mo- ment in the frontal plane is in the middle graph (B) and the torsional moment in the transverse plane is in the bottom graph (C). The gray bar indicates the fixation strength of ce- mentless implants (14). Reprinted ~vith permission from Bergmann G., Graichen, F., & Roh/mann, A (1995). Is staircase walking a risk for the fixation of hip implants? j Biomechanics, 28(5), 535-553.
100 vantage of the ~lhductors (Fig. 16-3) (Delp, \":'om;:tttu, &. \\Vi:\"son, 1994). Joint contact forces therefore E 020\" extension 0.3 ~ should be minimized with sm~t11 1H..'ad-neck all!!les [J] 20\" flexion (Johnston, Brand, &. Crownillshicld, 1(79). 'De- .3 .34\" flexion 3 cre~\\sed heml-IH_'ck angles also improve joint stability Cw 70 through increased congrucnce hv tUI-ning the 1m 68° flexion go k'moral head del.~pt..'r into Iht..' aCt..'tabulul1l. j\\/toving E the greater !rochalllL'r b{crall~' also increases the ~ 60 ~ mechanical ad\\'antagl..' of tilL' ahductors. Clinically. g'\". 50 0.2 :'\"D increased abdlll:lor/adducLOr stn:ngth has hC1..'1l asso- o icws 40 !i! ciated with incl\"easC'd neck length and a morc distal ~ 30 §f 0.1 \" greater trochantcr position (HuJ\"\\viti'. & i\\ndriacchi. E 20 1997). Transverse i'j\"j 10 Neck angle and neck length have an impact not olll~' on I he ~\\bductor muscle force and resultallt o hip force but also on the bellding rnomenls in the Resultant proximal felllur_ A varus hip or an illcrcasl.,d neck Translation length increases bending moments in the proximal Distal motions as a function of flexion angle (mean and standard deviation) There were significant differences be- felllur by inCrL'~ISillg. lill.' fllOllll.'nt arm or the forces tween 20 and 34° and between 34 and 68° for the resul- lransmitted along tile slwft or tilt.: felllur. Proslhc- tant translation and the transverse rotation, but for the sagittal rotation only the difference between the 20 and S('S IllUSl be designed to resiSl these bending mu- 34° was significant. Reprinted with permission from Berzins, A, Sumner. D.R., Andriacch!; T.P, et ell. (1993). Stem curvature ments. Decreasing the neck k'ngth or increasing and toad angle influence the initial relative bone implant motion !he ht.~ad-neck angle (yalgus) will decrcase the of cemenrfess femoral srems. J Orthop Res. 11 (5), 758-769. bending 1l10mcnts in the steIn; howcver, these al- • terations compromise thc abductor function and incrcas(' the joi nl react ion rorc(~ (Case 5t ud.v 16-1). tion produced by out-of-plane loads when the hip The abductors art.' important during the singlc- reaches greater ranges of Oexion. legged stnnec phase or gail because their contrac- tion pulls the rim or till' pclds toward thL' grt..'ater trochanter and prl..'\\1:nts tlK contralateral side of Reconstructed Joint Abductor Momenl GeOinetrv muscles arm Alterations in joint anatomy impact on hip biome- Varus Valgus chanics by altel-ing the contact area, the contact force, and the strength and moment-generating ca- The abductor mechanism changes with head-neck angle or pacity of the muscles. The effect of variations in ac- neck length. A valgus neck angle decreases the moment etabular position, anteversion angle. head-neck an- arm, whereas a varus neck angle or an increased neck gie. neck length. and joint center location on the length increases the moment arm. ReprinrecJ with permission muscles' capacity (moment-generating capacity) from Hun·vitz D.E. &- Andric1(chi, If {I 99B). Biomechanics of the and the resultant hip force have all been mathemat- 1111). In 1. Callaghan, A. Rosenberg. 8· H. Rubilsh (Eds.). Tile Adult ically modeled (Delp & Maloney, .1993; Doehring et Hip (pIJ. 75-861. New York: Raven Press. aI., 1996; Johnston, Brand, & Crowninshield, 1979). These analyses assume that alterations in joint and L------------- femoral geometry do not alter the manner in which subjects perform the activities. The mechanical ability of the abductors is af- fected by head-neck angle, neck length, and joint center position, all of which are frequently altered during a total hip replacement. A decreased head- neck angle (varus hip) increases the mechanical ad-
Cemented Total Hip Replacement A65-year-old woman, with chronic pain caused by as· teoarthrosis in the hip. Severe cartilage degeneration and reactive periarticular bony changes leading to loss of congru- ency had been affecting her hip function and daily activities such as gait. After a careful documentation of the patient's history, physical examination, and radiographic information, a total cemented total hip replacement is performed (Fig. (sI6·1·1). The picture shows the stage of the wlal hip arthroplasty af- ter surgery. The head-neck angle is in valgus. This valgus posi- tion will decrease bending moments in the stem prosthesis but will increase the joint reaction force as a result of reduction in the mechanical advantage of the abductor muscles (lever arm ; shortens Ie)). The figure is marked with the line of action of the abductor muscles (A), the line of gravity (while line), the line of ,the lever arm (b) from the center of rotation to the line of gravity, and the line of action of the reaction force at the esti- , mated point of contact of the femoral head and the acetabular unit (J). In addition, the line of gravity (white line) is depicted with the line of action of gravity force (W). The picture was made during a single-leg position, in which the line of gravity moves toward the hip, decreasing the abductor moment arm relative to the body weight. Thus, th!? patient adapts her gait and posture to decrease the de- mand on the abductor muscles by leaning over her leg and thereby limping a little. the pelvis from dropping as a result of the weight Although the models can predict hip joint center of the swinging leg and upper body. If changes in sitions that minimize the resultant force, these p the neck angle, neck length. and joint center posi- tions may not be practical for the acetabular co tion decrease the abductor moment arm relative to ponent. Pathological conditions strongly inOue the body weight and no compensation in [unction the potential locations of the hip cente!: For occurs, then an increase in the resultant force stance, osteoarthritis h-equently results in would be expected. Subjects may, however, adapt femoral head being displaced laterally, superio their gait to lessen the demand on these muscles and posteriorly. when the moment-generating capacity of the ab- ductors is compromised. Large joint forces are analytically predicted hip centers that are located superiOl~ lateral. Alterations in hip center location have a large ef- posterior as compared with the original locat fect on the moment·generating capacity of the mus- (Johnston, Brand, & Crowninshield. 1979). H cles and the resultant hip force (Doehring et aI., joint forces were obtained with superior latenll 1996; Johnston, Brand, & Crowninshield, 1979). Pre- centers in an experimental sel-up in which a load dicted joint forces are minimized when the joint fixture simulated the hip abductors, adductors, center is moved medial, inferior, and anterior (Fig. extensors during single-legged stance and s 16-4). This position maximizes the moment-generat- climbing (Gore et aI., 1977). Superior displaceme ing capacity of the abductors and brings the hip cen- of the hip center reduce the moment~generating ter closer to the line of action of the foot floor reac- pacity of the abductors, adductors, flexors, and tion force, thus decreasing the external moment that tensors as a result of alterations in muscle leng needs to be balanced by muscle forces (Doehring et and moment arms (Doehring et aI., 1996). Incre aI., 1996; Johnston, Brand, & Crowninshield, 1979). ing the neck length or advancing lhe gre ,./0.,0' C, _w·.'·
8 Joint Contact Force trudl;:IIlICl' pani~III.\\\" compt'llsal('S for lhl..'se lusses Stair in rI1u:-.ck m()lIll·llI-gt,\"lh.'r~lting l'aracitk-s (Delp, 6 KOlllattu. ~ \\\\'j.'\\:-'Oll, 19lJ.i), E In general. Ilk' ;:1il.. t1.\\,tical and \\..·.\\lk·'·in1l..'1H ..'t1 rc- slIlb on ,ill.' d\"fl..,\\.:t or juilll gL'(Jl1lclr.,\" {In hip ioint ·'w\" km.:cs al\\: l'OIlSisll'llt \\\\'ilh dinic;:l1 p;:\\liL:llt stu'dies. SO 4 CliniLal sllH.liL'S hau.: <.ISS(l(.:i;:\\kd ink-riu!' functional u~ 0 lIulcOnlt.::-- \\\\'jlh supcrior placl...·lllcl1l or till..' jnim CCn- IL'r (Box \"-\\..: ;'\\Clblc. 1093) and h~I\\'L' assoCi;:lled dc- fI) crt:;:tSCS ill abductor stn:llgtll and I(ISS of p;:lssive hip 2 ik'.\\iofl I'l'lOliOIl \\dlll sup1,.·rior pIaCt.:rlJCIlt:-. of the joint l'L'nlcrs l.llllL'sS COl1llk'llSall'd \\\\'illl ;:111 illlTtased neck k'll!.!lh (Hurwitz ~ i\\lldri;:\\cchi. 1997). Hiohcr fl..'n~()ral loo:-.cning ra1l..'s l1a\\'L' Ik'l.'ll ~lssoCi<llt'd ~\\'i{h joint cL'l'lll'rs placl.'d slqK'I'iur ~llld lalL'r-aI as opposed lu 1!1osI.· placed ill all anat(1Il1icd POSili(lll (\"'{odel' et al.. 1988). ,,-hill.' highl..'r \\'o!ulnl..'lric p(,lyelhyk'llc \\\\'t..'~l!' is :tssuciatl'd \\\\'itll dL'crl'~lsl'd k'llloral ollsds ~Hl(.l dClTL';\\SL,d ahductor 1'll1l1l1l'1'll i.11\"lllS (Robinson ('{ al.. llnpllhlishL'd d~ll~1. ThL' cllckl uf illlpblll posi· D lion in pul,h'th.\\'k'lll..' \\\\\\.'~Ir in IOl~11 hip anhroplasly). 0.4 Prosthesis Moment Sf eIII Po., if iOfl IIi fII ill flie Stair 0.3 r'e /110 m I Co iI {II zE Both clinical L'.'\\!kTit.:I1C!...' ;'lIl',1 hi()mL'ch~lllic~ll stn..'ss an~ll.,:ses :->UggL'St that a \\'alglls slt..'lll position pro- 1 ,'idl's hl.'tll.T rl..'slilts t1li.tll <.I(K'S ~l ,'ants position .c o(Alldri~\\l'chi 1.'1 ~I1,. 1076: Culli:-; & n.:,'l!.ull , 1977; ,~ 0.2 Gi.lli.Ulk'. RoslOh'r. I.\\: DO,de, 197)). Fillill..' I.'k'ml'nt SO !l1o(!L,ls indic;.I!I.' thai ,'i.II·llS slt..'llb sustain grl.'i.lll.'!' pL'i.lk sll\"l.'sscS lhall thosL' ill \\'algLls or llL'lllr~d for Ih..--: ~ same loading condiliol1s (.-\\lldrii.lccili l..'1 ;;11., 1476). SubjL'l'IS with :'1 .\\'hIL'IIL'I'\"cL'lllcnlL'd Stl..'lll ill ~l ,'antS \"o positi()l1 \\\\'~llk \\\\'ilb 11101\\.' ~d)[lorl\"ll~l1itics thall those \\\\'ilh a \\'alglls sll'l'll tiL-spilL' thL' I':ll'l lhal :-'lIbjL'Cl~ in fI) hoth groups h~H.1 an L'xcdlL'IH clinical OlllL'(IIllC (Hodg(', :\\ndrii.lcchi. & Gal~llllL', 19tJl). Tltl..' gail ab- 0.1 llorllli.llitiL,:-; in till.' \\';:II'lIS group may be sllggcsth'L' or <l bi{ll11l.'chaniGII ~ldapl~lli(Hl 10 I.'illle!· nlicromolioll ADADAD ur ahnol'lnal SlI\"L'S:-; patterns prL's\\.'nt in \\'i.I!\"lIS stl'llb. Peak contact forces at the hip and peak moments about Pcripmst!wlic BOi!e Loss the neck-stem junction of the prosthesis during level walk- ing, stair climbing, and rising from a chair under the fol- Pcriprosthetic bOlk' loss associ'lled wilh uncc- lowing conditions: A 0=0: normal location of the hip center Il'lclltcd rL'rnoral slerns has bL'l'll \\\\TIl dOClllllCl1\\l'd and 0 = hip center 20 mm medial, 20 mm inferior, and 10 (Br:-\"i.l11 t'! al.. 1994: Engll & Boh,n1, I 088; En~h L'l mm anterior. Peak forces were based on an optimization :.d., 1992L and concerns han:' lk'cn raisl.'d \\\\'itll rc- approach that minimized muscle stresses. Reprinted with gi.\\nls to the long·lerm clinical impliG\\linns or this c..permission from Johnston, R. Brand, R,A., & (rownins!lield, phL'nOllll.'llon. OSI('olysis, stress shielding, and gL'Il- R.D. (979). Reconstruction of (he hip. J Bone Joint Surg, 6IAIS),646. ',-:
eralized limb unloading all may play a role in Bobvn, 1988; Engh el al.. 1992). Au lOpS), stud ./0~; ;periprosthctic bolne loss\"lvvearl dcb,:islhas (been im- have also shown that the lower the bone mine ;,:,,::,'plicated in osten ysis ane imp ant lai lire Amstutz density or content 01\" the contralateral femur, \"i;, , et aI., 1992). Wear particles f!'Om the polyethylene greater the reductioll in periprosthetic bone on and other implant materials are found in the joint affected side, which further implies that preope Quid and adjacent tissues. Even small amounts of tive bone mineral density' influences the extent ~vear can generate large numbers of polyethylene postoperative bone loss (Engh et aI., 1992; Malon particles that are as small as or smallcl- than a mi- el aI., 1996). crometer.. These wear particles resuh in a foreign- ~'/ body· reaction with increased macrophage activity Forces at the Knee Joint and intercellular secretion of mediators that slimu- The knee joint depends primarily on soft tissucs \" late osteoclasts and I'esult in periprosthetic bone stability while sustaining large joint reaction for loss (Jasty et al.. 1993; Willert & Buchhorn, (993). at the tibiofemoral and patellofcllloral articulatio unexpectedly\" undersized or unstable compo- The peak forces during walking predicted by have a higher incidence of osteolytic lesions analytical modcls (Morrison, 1970; Schipplein than do those with stable fixation (Noble, 1995). Andriacchi. 1991; Seireg & Arvikar, 1975) val' Stress shielding results from a decrease in the from 3 to 7 bod~· weights and arc similar to th stress distribution in the fenloral bone as a result of measured in vivo (2.3 to 2.5 body weights) in single patient with a distal femoral replacement a the presence of the implant stem thal has a greater rotating knee hinge (Taylor et al.. 1997) (Tab 16-1 and 16-2). Thc magnitude and cyclic nature or equivalent mechanical stiffness as compared the compressive force in the tibiorelTlOnd joint with the femur. Changes in the loading environ- important considerations in the design of a to ment rcsuh in bone remodeling. Once bone in- knce replaccment, General chnracteristics of growth occurs in the cemcnllcss prosthesis, load predicted forces at the knee joint show three pe transfer can OCClIr through these areas of bony at- tachment. However, bone remodeling docs not re- Latera! Medial sult in the restoration of normal cortical strain lev- % change % change els (Engh et aI., (992). The fit of the prosthesis within Ihe femoral canal (JaslY et aI., 1994). as well Percent change in the strain energy density within the m as the material properties (stiffness) of the stem dial and lateral regions of the medial and lateral cortex f (Cheal, Spector. & Hayes, 1992; Weinans, Huiskes, a three-dimensional finite element model. The greatest r & Grootenboe,; 1992), affects the amount of stress- duction in strain energy density was found for the femur shielding (Fig. 16-5). with the stiffest implant (CoCr). The femur with the most flexible implant (Composite 1) had the smallest reduction Bone loss also can result from limb disuse strain energy density (Camp I and Comp2. composite mat (Goethgen et aI., 1991). Postoperative subjccls with als; Cocr, cobalt chromium; Ti. titanium) (Personal commun total hip replacements continue to walk aSy'lllmctri- tion. RN NariJrajan). cally, with decreased forces on the operated side as compared with the contralateral side (Bl)'an et aI., • -. 1996; Long et al.. 1993). Tibial bone loss is unaf- fected by implant characteristics and most likely re- sults from generalized limb disuse associated with asymmetries in joint loading conditions. A 160/0 de- crease in proximal tibial bone n1ineral content has been demonstrated in subjects with long-term total hip replacements, and this decrease has been re- lated to an as)'mmetry in the peak vertical interseg- mental knee force during gait (Bryan et aI., 1996). Preoperative gait mechanics of patients with hip OS~ teoarthritis have been shown to be correlated with the bone mineral density of the proximol femur- (Fig. 16-6) (Hurwitz el aI., 1997). The greater the bone loss preoperativelv, the less stilT the femur and the more that likely stress shielding and associated bone resol-ption will occur postopcratively (Engh &
during stance phase. Thus, failure as a result of rolling of the femur on the tibia as thL' knel' flexe cyclic fatigue of both interfaces and implant materi- two·dimcnsionall1lodelusing input from gait an als is an important consideration in lotal knee re- sis has demonstrated the presence and location placement designs (Ducheyne. Kagan. & Lacey. tractive forces on the tibial surface (\\.ViIII l1ler & 1978; Kagan. 1977; Landy & Walkec 1988). In acldi- driacchi. 1997). The model shows that a \"\"versa tion. the portion of the tibial plateau that is loaded the tractive force occurred at the postcrior end varies with knee flexion angle (Fig. 16-7). In many the contact region, suggesting that the effect 01\" t designs of total knee replacements, a large contact live forces should be considered in the evaluatio area only occurs for a limited portion of the knee damage mechanisms. Different types of gait range of motion with a I1luch smaller contact area cause different tractive forces on the tibial surfa occurring at other flexion angles. The smaller the Thus, one of the importanl variables in the con contact area, the larger the contact stresses. eration of factors leading (0 POI~:L·th.\\·lenc d~lm may be a variation in the tractive force associa One consideration in the Factors producing wear with the gait of a particular patient. on the tibioFemoral articulation is the tractive '\"'in •• • •• • • c 2 • w •• • 0 0 •• 2 2 rn -2 0 • •• <;; -2 c • 1.0 0 234 5 ~ Adduction Moment cw (o{;, body weight· height) 0 lD g C ~'\" 0 .e\"... <;; iii i\" OJ .'\"-w0 • rn -2 E 0 Z 5 10 15 20 25 30 0.0 0.2 0.4 0.6 0.8 Dynamic Hip Range of Motion External Rotation Moment (degrees) (% body weight· height) The hip range of motion and external adduction and exter- lady, the anterior fibers of the gluteus medius and minim nal rotation moments were significant predictors of the nor- are recognized as primary internal rotators. Thus, the de- malized bone mineral density of the greater trochanter in creased external rotation moment in early stance may als the patients with hip osteoarthritis. The abductors are the reflective of decreased abductor muscle forces. Modifi~ w primary structures responsible for balancing the adduction permission from Hurwitz. D.E.• Foucher. K.C Sumner: D.R.. et moment. Because the abductors insert on the greater (1998). Hip motion dnd moments dtJring gail relMe clireCily 10 trochanter, a reduced adduction moment may reflect re- proximal femoral bone mineral density in pc1tieflls \"''/Ith osteod duced forces in this region and may result in bone loss. Simi- tis. J Biomechanics 3 I (lOj, 919-925
Posterior Adduction Moment During Gait 60 and Medial - Lateral Joint Loading Laleral 70~:'50 .70~95g0 Medial Knee Adduction Moment ROIII~ 20 o 6020 _'0 Roll I- <.Dt 10 0+-1--------. 40 30 20 10 0 10 20 30 40 em c c ~\"E ~ 20 The tibial-femoral contact moves posteriorly with knee 40 flexion. The contact on the lateral side moves posteriorly much more during flexion (0-20°) than does the medial 1-----1 side because the lateral femoral condyle is rolling on a Heel Slance T larger radius than is the medial femoral condyle. Beyond 20°, sliding motion begins on both condyles. Reprinted with strike phase o permission from Andriacchi, T.P., Sran...vyck, IS., & Galante, J.O (J 986). Knee biomechanics and (Oral knee rep/~Kement. ArttlrO- plasty, 1(3),211-219. Medial-Lateral Load The force on the medial Distribution compartment increases as the adduction increases One of the major problems limiting the use of total knee replacement in the 1970s was tibial compo- An illustration of the adduction moment during walking nent loosening (Bargren, Blaha, & Freeman, 1983; and the resultant greater load across the medial compar Ducheyne, Kagan, & Lacey, 1978; Hsieh & Walker, ment of the knee compared with the lateral compartme 1972; Kagan, 1977). Tibial component loosening Reprinted wirh permission from Andriacchi, T.P. {1993J. Func- was related to the load imbalance between the me- tional anafysis of pre- and post·knee surgery: TOlal knee arth dial and lateral surface of the tibia that often occurs pfasty and ACL reconstruction. J Biomechanics En9, 115. during walking and other activities. The early de- 575-581. signs of the all-polyethylene tibial component were not sufficient to sustain this load imbalance. Our-ing between the proximal medial and lateral ti walking, approximately 70% of the load across the (Hurwitz el aI., 1997). knee joint is normally sustained by the medial com- partment of the knee (Morrison, 1970; Schipplein & The load asymmetry at the knee resulting fr Andr-iacchi, 1991; Han'ington, 1983), The adduction the adduction moment during'· gait provides moment is a major determinate of the load distrib- strong rationale for maintaining proper limb alig ution between the medial and lateral plateaus (Schipplein & Andriacchi. 1991) (Fig. 16-8). In fact, among normal subjects, the adduction moment is a significant predictor of the bone dislribution
,I Mechanical Knee Arthroplasty axis of the ,I femur A70·year-o:d man ~tl;ierlng 110m dlSabhng leit knee paUl as a result 01 (1 S(''lo?re genu '!arUS ane! pr()(JresSlve I left knee JOint degeneration The- Clbnonr:alload resulting from the g€'nu varus ddormlty erc-ates ii load unbalancE:' I characterl.zed by a decrease in conlan <w~a al llie I,lleral I left tibia plah~au dncJ C'Hl !nm'ii~;C' of the contdCl SlrbSes '::1[ I Femoral the medial tlbla plateau resulting III progn~sslve ',,'edr in the shaft axis Mechanical I Femoral medial tJblofernoral compartrncn:. (Fig. C~ i 6·2· i). Tibial shaft axis The patient hciS an H11palfll1e-nt and is lJn,1bJe to ':\"Jlk axis of the I shaft axis I Mechanical more than half a Imle He '.·\",shes to m,llfHatll hiS (1CtI'lE' lower I axis althe lifestyle. A first trial of (onservaUve tre,l;rnent was unsuc· extremity I tibia (tibial cessiul. A careful examination ensures ;he presence of a shaft axis) functiondl extensor nWel1dnlSrn, !v1ordwer. COl'11p!QtE' I imaging sIUda?'> confirm the 5e\\h'.:'I'(~ ,:!rtICuliH SUrf,'1(\\:' cie- I orgenera !Ion. The de-Clsion kne-e Mthroplas!y \\\\''::$ made I to avoid il more ')('vere JOlllt (:e-gel10fdtion an(l unprove patient\"s liiestyie I I I ,r I Left, In a lower extremity with varus deformity, the me· chanical axis passes medial to the knee. Right, When align- ment is normal, the mechanical axis of the femur is in line with the mechanical axis of the tibia (tibial shah axis). The line represented by the mechanical axes of the femur and the tibia is coincident with the mechanical axis of the lower extremity in this situation. Reprimed with permission from Kracholl'.t. K.A. (995). Surgical principles in total knee arrhroplasry: Alignment, deformity, approaches and bone cuts In J.j, Callagclfl. D.A. Dennis, W G. Paprosky, er at. (Eels';' Or- thopaedic Knowledge Update: Hip and Knee Reconstruction (pp, 269-216). Rosemont, IL: MOS. • ment following total knee replacement and uni R Case Study Figure 16-2-1. compartmental arthroplasty (Fig. 16-9) (Bargren, Blaha. & Freeman, 1983; Brugioni, Andriacchi, & Galante, 1990; Weinstein, Andriacchi, & Galante, t986). Knees with a vanls alignment are more likely to have a substantial load imbalance that cre- ates stresses that could eventually lead to tibial component loosening (Case Study 16-2). ,Increased wear has been demonstrated in the medial com- partment of knees initially in varus preoperatively and in the latet'al compartment for those in valgus preoperativelv (Wasielewski et aI., 1994). To ad- dress problems associated with this load imbal· -'l\"~4.~.
ance, tibial component designs were modif-lcd to *100 Abnormal have a mewl backing of the polyeth~'lcnc articulat- pattern i.m!: surface, and surgical instrumentation was rnod- ifi~d to allow for proper alignment or the mechani- cal axis. tello{erl1oral Joint o Anatomical troc Loads 5.0 the incidence of tibial component loosening sub- stantially decreased in the eady 1980s, patello- -c f 2.5 Non-anatomicallro fcmoral problems began to emerge as a primary is- E\"\" s: sue in total knee replacement (ROI'abeck et aI., 1995; OlD Rosenberg et aL. 1988), The magnitude of the rell'o- :i: ~ 0.0 --I''-------->,--~f_-1''''---..-....= - - - patellar force as well as the contact area on the retropatellar surface varies with knee Oexion angle 2,5 (Ahmed, Burke, & l-hder, 1987; Hubeni & Haves, Swing 1984), Loads between two and three body weights have been reponed at the putcllofcmoral joint for (Top) The moment tending to flex the knee during , various activities of daily living. It was shown that climbing demonstrated differences in stair climbing for walking, the loads arc relatively' low but because tion between the patients with nonanatomical fem the knee flexes beyond approximately 40\" during ac- trochlea and anatomical femoral trochlea designs. tivities such as stair climbing, the loads can reach Reprinted with permission from Andriacchi, T.P., Yoder. D levels of several body \\veights. The high magnitude Conley, A., er 8/. (997). Parel/ofemaral design influences of these forces poses a great risk lO the mechanical tion following coral knee iJrihroplasty. j Anhroplasty. 12( integl\"ity of the implant as well as the implant's fixa- 243-249. (Bottom) Patients with the nonanatomical tion to bonc. femoral trochlea had increased knee flexion in late phase (shown by It). The increased knee flexion was The external knee nexion moment as measured anced by an increase in the flexion moment (balanc with gait anal~'sis is reflective of net quadriceps net quadriceps contraction) during late stance phase muscle activity. Alterations in the external llexion Reprinted with permission from AndriacclJi, T.P. & Hurvvit momenl during stair climbing have been related to (1997). Gail biomechanics and total knee arrhroplasly. A both the curvature of the patellar nange (Anclriac- Knee Surg. 10(4), 255-260. chi & Hunvitz, 1997) and the height of the patella (Mat,tel et aI., 1990), Two groups of patients with the increase in knee flexion. Patients in the different total knee replacement designs that group did not have these abnormal charac differed primarily in the anterior curvature of tics. In the design or a total knee replacemen the patellar flange in the region or the femoral replication of an anatomical curvature o trochlea \\vere evaluated \\vhile climbing stairs. femoral lI'ochlea requires more bone resecti Both groups were selected on the basis of a good the distal anterior femur than does a nonana clinical outcome. One design group (nonanaloml- cal curvature. Thus, many designs compro cal trochlea) had a smaller radius on the patellar between the restoration of normal anatomy flange that caused the patella to articulate more more extensive reseclion of the distal femur. anteriorly and distally than the second design results suggest that replication of normal pa (anatomical trochlea) that had a larger radius in femoral anatomy is important because it will this region. The second design more closely repli- 4\\ significant effect on both function and pa cated the femoral trochlear anatomy. The group loading. with the design that placed the patella more ante- riorly had a slight buckling (increased Ilexion) of the knee in late stance (Fig, 16-10), There was also an increase in the moment at the knee sustained by the quadriceps in late stance concurrent with • ~•.. <
Joint Line Height Interaction Between Tibiofemoral Joint Line Position and In some instances, the joint line is elevated following total knee replacement, \\vhich results in an inferior Patellotemoral Mechanics movement of the position of the rctropatellar con~ tact on the femur: Elevated joint lines, which affect Slair Climbing Knee Flexion Moment patella function and patella subluxation, have been and correlated with wear patterns (vVasielcwski et al., I994), decreased longevity of the patellar component Joint Line Change (Rosenberg et aI., 1988), decreased clinical outcome (Junnosuke et aI., 1993), and decreased range of mo- . .. .f - - - - - - - - . •/'\"\"-\"\"\":......----- tion (Rittel~ Faris, & Keating, 1988). The height of /' the patella with respect to the joint line has been as- ••A sociated with signilkant changes in the flexion mo- ment of the knee during stair climbing (Fig. 16-11). · At-- - - . • L-' • vVhen the patellar position was changed IS mm in- ferior to its normal position, more than 50% reduc- 15 10 5 a tion in the flexion mornent during stair climbing was noted (Martell et aI., 1990). Change in Patellar Contact (mm) Posterior Cl'llciate Ligament The position of the retropatellar contact is related to the Differing opinions exist with regards to the benefits location of the joint line. Typically, in total knee replace- of retaining, substituting for, or removing the pos- ment the joint line is shifted superiorly, causing a down- terior cruciate ligament (Andriacchi & Galante, ward movement of the location of the retropatellar con- 1988; Li et aI., 1995; Rorabcck et '11.,1993), with no tact at any flexion angle. The retropatellar position is definitive answers with regard to which design re- related to the flexion moment during stair climbing. sults in better wear rates, clinical outcome, or Reprinted Wiih permission from Andriacchi, TP (i 993), Fune- range of motion. Sonle studies indicate no substan- lional analysis of ,0((-: and !Josi-knee Sur9('ry: Ibldl knee arthro- tial differences in loosening or radiolucent lines plasty and ACL reconstruction J BiomechilflicS Eng. 115. (Ranawat et aL, 1993; Stern & Insall, 1992; Wright 575--58 , & Bartel, 1986) or range of motion (Hirsch, Lotke, & MOlTison, 1994) between the different designs. In component to increase the contact area and mll1l- fact, the preoperative range of motion may have a mize the stresses and resulting polyethylene wear. As greater influence on the postoperative range of mo- discussed earlier, the lateral side of the knee can be tion than any influence attributable to design dif- unweiglHed during the stance phase of gait and can ferences (Han'cy et 'II., 1993; Maloney & Schur- result in the knee being in a V:'lrttS position. The flat man, 1992). The decision of whether to retain or articular surfaces frequenliy used with posteric)1\" c\\·u- remove the posterior cruciate ligament imposes de- ciate*l\"etaining designs permillarge contact areas but sign criteria on the amount of constraint that the articulating surfaces need to provide. Retention of the posterior cruciate ligament re- quires that the joint line be accurately reproduced for the kinematics to remain normal. Othenvise, the knee kinematics can be altered and loads across the joint may even be increased. Limited knee flexion (Rillcl; Faris, & Keating, 1988) and posterior poly- ethylene wear (Swany & Scoll, 1993) have both bcen associated with posterior cruciate ligaments that are too tight and result in posterior tracking of the femur on the tibia. Posterior cruciate-retaining designs gen- erally have an unconstrained but conforming tibial
can result in edge loading and high stresses if the im~ Stair Climbing in Patients Following Total Knee Replacement plant is tilted. Observation Interpretation Removal of the posterior cruciate ligament neces~ sitates that posterior stability no\\v results from in~ I. 10.0 PCl retained Normal roll creased tibial-femoral constraint. In the absence of a Stance back increases posterior cruciate ligament, the posteriorly directed ~ quadriceps leve shear force is instead sustained by the interfaces of arm (/) with fle the articulating surfaces (Andriacchi & Galante, ~ 4.0 1988). Although more constrained designs are inher- ently more stable, they may limit the passive range C of motion by restricting the amount of rollback needed lo achieve deep knee flexion (Fig. 16-12). §'\" 0.0 Substituting designs are also more constraining with a cam mechanism used to maintain the posterior po- ;:;; sition of the femur relative to the tibia. PCl Variations in the knee kinematics and moments during stair climbing suggest that posterior cruci- removed ate~retaining patients have more normal function \\vhile ascending stairs than do patients with knees PCl/ ¢~~~ PCl removed in \\vhich the posterior cruciate ligament is removed back constrain (001'1' et 'II., 1988; Kelman el 'II., 1989). Patients with retained and leverarm cruciate~sacrificing knee replacements had a ten- reduced dency to reduce the moment sustained by the quadriceps by leaning forward during the portion of Patients with PCl the support phase of ascending stairs when the removed tend to lean quadriceps moment would reach a peak value (Fig. 16-13) (Andriacchi, Galante, & Fermier, 1982). This forward reducing the finding was also consistent with a study in which net quadriceps moment electro111yographic activity' was measured during Patients with designs that remove or substitute for the 15,----------------, posterior cruciate ligament tend to reduce the knee fle ion moment and thus the resulting demand on the qua ceps. The mechanism they use for the adaptation is a fo ward lean of the torso. The biomechanical explanation the adaptation seen in these patients can be related to normal posterior movement of the tibiofemoral contac with flexion, which is reduced when the posterior cruci ligament is removed and constraint is added to the arti lar surface. Reprinted with permission from Andriacchi, TP (1993). Functional analysis of pre- and post-knee surgery.' To knee arthroplasty and ACL Reconstruction. J Biomechanics E 115.575-587. OLL_ _.LJ_ _--'--L_ _l-L_ _-LJ stair climbing (0011' et 'II., 1988). In that study, tients \\vith posterior cruciate ligament-sacrific Normal Cruciate Cruciate Cruciate designs required increased use of the soleus mus knee retaining excising substituting while stair climbing. The increased soleus acti prosthesis prosthesis prosthesis was suggestive of a forward lean in the patients w posterior cruciate Iiga111ent-sacrificing designs si1 Femoral rollback as a function of prosthesis type during I'll' to thal described in other sludies (Andriac stair descent at 900 of knee flexion, Reprinted with permis- Galante, & Fermiel~ 1982). In a 1110re recent stu sion from Mahoney- O.M., Noble, pc., Rhoads, D.O., et al. by Wilson et al. (1996), patients with posterior C (994). Posterior cruciate function fof/owing total knee arthro- ciate ligament-sacrificing designs were compa plasty A biomechanical study. J Arthroplasty, 9(6),569·-78. with normal subjects while walking and stair clim ing. Although the study did not report a statist difference, there was a 25% reduction in the p
TOlal Knee Replacement radius of the tibial component increases relative to Frontal Plane Profiles that of Lhe femoral component, {he conformity de- and Contact Sfress crcHSCS and the contact stn:ss between lhe two components inc]'e~\\scs (Fig. 16-14). Conformity is 4,0..----------------, frequently' used 10 characlcrizc the articulation be- twcen a dished tibial surface and a rounded r~o 3° (,,,\" fL'Jlloral component, buL it can also be used [0 de- 1 .iii 2.0 scribe the articulation between a flat tibial surface ...•.. and a nal femoral component. A flat geomctl'y has 8uc'\" 1.0 a conformit.v of one and docs nOl constrain lhc 1'0 , lational or translational 1110\\,el11ent as required with a posterior crudate-retaining design. A Radius Ratio R2/Rl dished ge0I11ctr.v achieves conformity and con- straiI1l as needed for the posterior cruciate liga- The predicted contact stress increases as the conformity menl-sacrificing designs. To maintain the sagittal decreases. Modified with permiS5ion from Andriacchi, T.P. & plane kinematics, the femoral componelll n111st Natardjrln, R.N. (993), Conformity and polyelnyfene damage in have a smaller posterior radius and a larger distal tOMI knee replacemenc (Internal Communication 97-51/0-206 radius. Thus, it is not possible to achic\\'c confor- 3Ml), Zimmer, Inc mil y' in the sagittal plane in bOlh nexion and ex~ tension (Fig. 16-15). • nexion moment during stair climbing between nor- mal subjects and patients following LOlal knee re- \"~iUS ralio extension placemenl. The lack of a statistically significant dif- ference in that study may have been associated with ~ the small sample size or the test population, The functional differences between the postel'ior cn-'Ciate ligament-retaining and sacrificing designs were associated with the normal posterior movemenl of the femur on the tibia (rollback) with flexion, This finding has been explained by the dynamic interac- tion belween the posterior crudate ligament and tibiofemoral rollback with flexion. The lever arm of the quadriceps (Andriacchi, Stanwyck, & Galante, 1986) normally increases with knee nexion. The func- tional adaptations seen among patients with cruci- ate-sacrificing designs were likely associated with the need to compensate for the lack of normal femoral rollback in knees in which the cruciale ligament is re- moved, Thus, rollback must occur in the early phases of flexion to have an appropriate quadriceps lever The radius of the femoral component varies with knee arm to sustain normal stair climbing. flexion angle. Thus, in the sagittal plane it is not possible Conformity to achieve optimal conformity in both flexion and exten- sion. Reprinted with permission from Andriacchi, T.P. & Nararajan, R.N. (/993). Conformity allcl polyethylene damage The degree of conformity between the femoral in lolal knee replacement (lnrerndJ Communication component and the tibial component depends on 97-5110-206 3MZ). Zimmer, Inc. •the ratio or the radii of the two components. As the
Section through metal femoral component DefeCIS Time Seclion Ihrough in plastic plastic libial componen1 Cracks propagating from defects Surface wear ~ and deforma1ion smallr::~r-:~~~=t--I~~::~~~\"\"\"'7 contacl Level of maximum shear area and sUess moves downward high slress as surface wear proceeds A model of delamination wear in polyethylene with intergranular defects. The lower dia- gram explains a possible reason why delamination does not occur if there are no defects in the material. Reprinred with permission from Wdlker. AS. & Blunn, G.W (997). Keynote Lecture 1/: Modem design of tocal knee replacement. In S NiwiJ. X. Yoshino, M. Kurosdka. K. Shino, S. Ya- mamoto (Eds.). Reconstruction oi ,he Knee Joint (p.o. 129-142). Tokyo: Springer· Verlag. • Constraint body particles (cement), and areas of high cont stress (Bargren, Blaha, & Freeman, 1983). Defe In general, proper soft tissue balance is necessary for in the polyethylene serve as sites for crack initiat a satisfactory outcome. In the presence of con;plex and propagation (Fig. 16-16). Increasing the po ethylene thickness to at least 8to 10 mm will red sbaotni~yfadcetfOOl)l-'mstitaibesilistuycihs aosftaenseuvneorbetvaianlg:,ubsledeafnodrmcoit~v~ contact stresses and should decrease the amoun wear (Collier et aI., 1991). Metal backin2 was straint must be provided by the implant design. His- tially introduced to more evenIv distribut-;; the lo torically, a hinged prosthesis, which provides maxi- to the underlying cancellous b~ne of the proxim mum constraint. was advocated for those patients libia. Howcvc,', overall component thickness need with severe valgus deformities. Hinged prostheses, to be increased following the introduction of me however, have been associated with a poor clinical backing so that an adequate polyethylene thickn outcome (Bargar, Cracchiolo, & Amstutz, 1980; Bui was maintained. The multivariate nature of & Fitzgerald, 1980; Rand, Chao, & Stauffer, 1987). A causes of pol:vethylene damage makes it difficult more contemporary alternative to the hinged pros- draw definitive conclusions on failure mechanis thesis is a highly constrained but unlinked total knee from retrieval studies. Moreovec the diverse nat replacement (constrained condylar knee). This de- of the contact stress distribution (compressive, t sign eliminates the need for li0o\"ament 0balancin.a bv sile. and shear) may result in differenl dama modes. Deformation, pitting, cracking. and ab gpurolavridipnoglymetehdyilaelnleatetirbailalstapboilsitt.-y Tthhriosu\"ahhigahlyrecctoann-- sion are all examples of polyethylene dama modes. strained design has provided adequate outcome for elderly low-demand patients with severe valgus de- Anterior Cruciate Ligament formities (Bullek, Scuderi, & Insall, 1996). During total knee replacement, the anterior crud -Polyethylene ligament is either already absent ~r usually , moved. In the case when only oTie compartment the knee requires replacement (unicompartmen Polyethylene damage has been closely related to its thickness (Collier et aI., 1991; Wright & Bartel, 1986), the material properties, the presence of third- .,
knee arthroplasty), both cruciate ligaments are fre- :\\Ildriacchi, TP. (1993). FUllclion;t! all;t1.\\'sis or prL'- ;\\Ild POst quentl.\\' retained. Gait studies of total knee replace- knee surgL'lS: TOI;t! kllL'L' ,lrthropl,lStv ,tlld :\\CL I'econstruc_ ment patients (Andriacchi, Galante, & FermieI~ lion. J Diollll'c!IIiJlic.\" Lug, 115,575-5:)1. 1982; Chao, Laughman, & StauffeI~ 1980; Kelman et aI., 1989; Simon et aI., 1983) show that normal func- :\\ndri:tcchi, T.P, 6.' C;\\l'lnte. J.O. (19t:t:J. Relcnliull (ll' IhL' DOS tion is not achieved in the majority' of the patients tl'rior nlll.:i;ill' lig,lIl1enl in tul;1I knee arthropl;\\SI~'. J despite improvements in stride lengths and knee Arthroplasty, SI3-,19. motion and clinical status. Numerous subsequent studies demonstrated differences in the flexion- .-\\ndri'lcchi. TP.. Gabnll', J.O., Iklvlschko, TIL l't ,d. (1976) extension moment during level \\valking. These vari- ,.\\ stress an;dysis 01' rl'[llOI';t! stelTI in lotal hip prustheses. J ations in the flexion-extension moments at the knee HOlle Joilll \"ill I)!\" 58:\\(Sl,61,s\",62·1. could not be related to an}' specific design but in- stead suggested the possibility that loss of joint pro- Andrial'clli, TP., G;danlc. J.O .. 6.' FLTlllil'l', R.\\V. (1982). The prioception or the functionalityl of the anterior crll- ciate ligament may have been a factor (\\J\\'einstein, inl'hlL'ncl' or loul! knee rL'pl;tCeIllL'lll design on walking and Anclriacchi, & Galante, 1986). The removal of the sl;\\ir-clilllhing. .I !lout' .loiul Sill)!\" 6-1:1, 132S. anterior cruciate ligament was common to all total knee replacement designs and may be one of the Alldri;\\cchi. TP. & Hurwit/, D,E. (1'107J. G;lil hiomeckmics factors that limit a patient's recover)' of total normal function. Similar patterns of abnormal flexion- ;llld tillal knee arthruplasl,v. ·1!11 j I\\I/(l' Sill',:!\" 10(4) extension moments (Berchuck et aI., 1990) have 255-260. been reported in patients foll()\\ving injury to the an- .'\\ndriacchi. T.P. & N:\\laraj~ln, R.\\:. (1993). Conl'ormil\\' ~ll1 terior cruciate ligament. polyelh,\\'lt:!lL' d;lIn;l):!l' in lutal knee replacelllelll (Interna Communication 97··S110-206 3\\lZ). ZillllllL'l'. Inc. Summar)! :\\ndriacchi, TP.. Stal1\\\\',vck, T.S., & Galanle. J.D. (1%6). Kne~ biolllech;lllics and tOlal knee repl;lct:JllL'llt. :Irlhro/dusty t The effect of forces on the stability of a total /(3),211-219. joint replacement depends not only on its magni- :\\ndriacclli, T.P\" l·odl'r. D.. Conle\\'. .'\\ .. L't ~i1. (1997) tude but also on its orientation and point of appli- Patelln!'ell1or;1! d('sign influL'nces funclion I'ollowing tota cation. knee 'lJ'lhroplaSI,\\'. .I :lohruplosty. 12(3). 2.L',·2-l9. Barg;II', \\V.L., Cracchiolo . .'\\., L\\.: :\\mstllll. H.e. I: 19801. Rt:sult Magnitude, orientation, and point of applica- \\\\'ith constr~tlned IOlal knee proslhl'sis in lreating se\\'L'l'('!Y tion of the loads at joints influence the stresses, dis:thkd p,lliL'nb ;\\Jld p:ltil'lItS \\\\illl 1',\\iIL'd tot,li kIll:\":' 1\\> bending ITIOments, and rotational moments of the pl,tce!1lent, J BOJlt, Joint Surg. 62:\\. 50..\\-\"512. implant, and are critical for the implant stability Ibrgren . .I.H\" Blah;!. .l.D., & Fret:rn;lIl. \\1.:\\,R. (1983). .'\\li~!Il and longevity. lTlL'nt in IOl,1I klll'e i\\rthropl:lst.\\', COlTL,I;ill'd ,tnd clinical ob S(·l'\\'iltions. Clin On!zO!I, /73. 17S, 3 Understanding the d.ynamic loads during daily' lkrchuck, \\1.. :\\ndriacchi, TP., Bach, B.R. Jr.. el ;d. (1990) activities provides critical information for address- Gait ;ld;tptalions hv p;\\li(:nls \\\\'110 ha\\'<.-' :1 deficienl :\\CI.,. ing clinical problems such as ITIechanicalloosening BOlle Joini Sur,!.:. 72:1. 871-877. of implants, amount of weac bone resorption, and Ikrgmann, G., Graichen. F., L\\: Rohlmann, .'\\. (1093J. !-lip join the choice of rehabilitation and surgical protocols. k':lding during \\\\',lIking ;lnd running measurwl in t\\\\'o pa 4 The evolution of total joint replacement has lil'nls. J BirJlncc!wuics, 26(8), 960 .·990. been aided by information generated from biome- chanical studies. Ikrgnwnn, G.. Gr;lichell, F., & Rohlmann. A. (199S). Is slair case walking a risk ror the fi\\;\\lion or hip irnpl;lI11s?.1 Bin REFERENCES lI/c·c!liIuic.,-, 28(31. 335-333. Ahmed, A.M., Burke, D.L., & Hvc!l-r, A. (1987). Force analvsis Berzins. :\\., Sumner. D.R .. & Andriacchi. TP. (19931. Stem of the patd!;:tr mechanism. j Orr/lOp Uus, 5, 69. . 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(1998 DOlT, L.D., Ochsner, J.1.., Gronley, J., et al. (1988). Functional motion and moments during gait relate directly to comparison of posterior cruciate·retained versus cruciate- sacrificed total knee arthroplasty. Clill Or/hop, 236, 36. mal femur bone mineral density in patients with h Ducheyne, P., Kagan, A., & Lacey, J.A. (19i8). Failure of total teoarthritis. J BioIllCC!Ulllics, 3 I( 10), 919-925. knee arthroplasty due to loosening and deformation of the Hurwitz, D.E., Guyton, J.A., Andriacchi, T.P., et a!. ( tibinl component. J 130111.' Joi/I! SlIrg, 60A, 384. I-low can,d fill in uncemented total hip replacemen Engh, CA. &: Bobyn, J.D. (1988). The influence of skill size fects the biomechanics of gait. 38th ..lIllI/wI ,Heetillg OrillOpaedic Rcsearch Society, I 7( I), 281. and extent of porous coating on femoral bone resorption Hurwitz, D.E., Stunner, D.R., Andri;:lCchi, T.P., et al. ( after primary cernelltlcss hip arthroplasty. CORR, 231, Dynamic knee loads during gait predict proximal 7~28. bont' distribution. J Biolllec!lallics, 31(5),423-30. Jast)\", 7vI:, Bragon, C., Lee, K., et al. (1993). Wear of po Engh, CA., McGovern, T.F., Bobyn, J.D., et al. (1992). A quan- ~'lene in total joint <lrthroplasty. In B.f. Morrey (Ed.) logical ;\\I(/teri(/l (/I/(! J/echallical COllsidi!J\"(/tioIiS orJoil titative evaluation of periprosLhetic bone remodeling after placeillellt (pp. 103-117). New York: Raven Press Ltd cementless total hip arthroplasty J BOlle Joint SlIrg, NA, Jasty, M., O'Connor, D.O., Henshaw, R.M .. ~t al. (1994). 1009-1020. the uncemented femora! component and the use of c Engh, CA., O'Connor, D., Jasty, M., et al. (1992). Otl<lntifica- influence the strain Lransfer to the femol',d cortex. tion of implant micromotion, strain shielding and bone re· thop Res, 12, 648-656. sorption with porous-coated anatomical medullary locking femoral prosLheses. CORR, 285, 13-29.
Johnston, R.C., Brand, R.A., &.: Crowninshield, R.D. (1979). illltllL'di~lIL' ri,\\:llion wilh 1(t~Hling \\'L'nical. fl'nlUr' horizon Reconstruction of the hip. J BOlli! Joillt Surg, 61..1(5), tal. j Bi<I!II('clu/IIics, 2-l( 1),37··48. 639-652. Pilli~lr. [(,M., Li.'i:. L\\I ,& .\\l~llliatop(l\\dll:-', C. (1986). ObsL'rv< Junnosukc, R., Saito. S., Yamamoto, K., ct .. L (1993). F~u.:tors lioll:> (In t!h: dfect of 1ll(J\\·L·ml·nl Oil hl.lnl· ill~rfJ\\\\'lh int influencing the po~toper.i.\\Ii\\'e r\"Hlge of motion in IOI~d knee arthroplasty. Bull Hasp Joillt Di.'i, 53(3).35-·W. pnroll:-.slIrbcl·d imrl\"llH~. C/ill O,.t/IO/I. 208. 10$-113. Kngan, ,\\. (1977). ~..kch'lnical C<lllS~S of loosening in k!lei,.· joint rcp1:\"lcements. J Biol1lecJulIIics, 10, 387. R~1I1:I\\\\'a!, C.S .. Fl,vl1n, \\\\\".F.. S:llldlt,r. 5., \\,:1 :I!. (1~93l. Long Kelman. G.J., Biden, E.N .. &. ()hil. et al. (1989). Gait labora- ll.'rm fesult:- fJf !Ill' tot:t! cOlld~'I:lr Klll'l' <Irlhrnpl:lst~·. (,li tory analysis of a poslerior cruciate-sparing total knee arthroplasty in stair ascent and descent. Cfill Orrhop, 248, OrihlJp. 286. 9.J-1 02. 21. Rand, J.A .. Chao. E.\\'.S .. ,,\\; St:lUffl.'r. R.W, (19$7). t\\illl..'lll:11 Kotz,lr, G.~·L, D;:l\\'Y. D:r., Goldbcrg, V.;\\,l., tl nl. (1991). rowtillg hillg<.: knl't.' ~lI·lhroplasty . .I !iOIl(' .Iuilll S/lI\"g 69. Trlcmch:rizcd in vivo hip joint fon:c data. ;\\ rcp0rl on 1\\\\'0 paticnts .Iftcr wt\"l hip surgery. J OrtJlOp 1?l.'.,', 9, 621-633. ~Sl)-497. Krnchow. K.A. (1993). Surgical principles in total knee i1rthroplasty: Alignmenl, deformity, appro:lches nnd bOlle Riltl,.·r. M,A., F;u·is. P.\\\\.. & KL·.i.\\lin~. E..\\\\. (198$). P()~ll'ri elliS. In J.J. Callag\"ln, D.A. Dennis. \\V.G. Paprosky. & A.G. Rosenberg (Eds.), 0I'111o/){/rdic KI/CJIl'led!,;e Update Hip (Iml (TIH:i;11e lig:1l1lclll b:dantill~ durin!,: Itllal krli.:l,.· ..\\nhro Kllee R(:(,.·OIlSIJ\"llClioll (PP. 269-276). RoseffiOll1. IL: AAOS. Landy, :\\'1.:\\'1. &: Walker, P.S. (1988). Wear of u!tl';:\\-high-molec· pl~ISt~·_ 1 ..lrlllf(lp!llsty. 3. 313. uhlr weighl polycthyknc components of 90 rctficved knee prosthl,.·scs. 1 Arthroplasty (SlIpplJ. 73-85. R(,raht.'ck. C.II .. Bournl·. R.B .. Ll,.·\\\\\"i:-. P,L.. d :11. (19<;3). Th Li. E.. Rincr. :\\1., :\\Ioilancil. T.. 1.'1 .d. (1993). Point-cOllnlcr- point \\Otnl km.'e arthroplasty. 1 Arthroplasty, 10(4), .\\lilkrfGalarlLl,.· klll'l' proslhl,.'si:- for I ilL' trl':\\IIllL'1l1 of o 560-570. Long, \\V.T.. DOlT. L.D .. Healy. B.. l..'l ,,!. (1993). Functional rl,.·- Il'o:lrlhrosis: ..\\ comparisoll of Ihl.: rL'slIhs of pal'li\"ll fix covcry of nonccmented total hip llnhroplast~'. CORR, 288. 73-77. tion with cellH:nt ..Inti fix:llinll without any Ct.·IlH:lll. 1 Bon ~'lahoney, 0.:\\>1., Noblt::, P.c., Rhoads, D.O., l..'l al. (1994). Pos- terior crllci<lte function following lotal knee ~\\rthroplasty. jO;IIT Sur;.:, 75. 402. A bionll.'chanicai study. J Arthroplasty. 9(6). 569-378. Maloney, \\V.1. &. Schunn<ln, OJ. (1992). The effecls or im- Ror;'d~L·d. C.H., DOlT.I.D .. Hofnwnn, \"\\.:\\ .. t.'l a!' (1995). Co planl design on range of nHHion after total knee anhro- tnl\\'l.:rsial issu('s in kn('L' ;ll\"lhropl:lst~·. Onhofl CI'(,I$,\"1I plasty. Total condylar versus posterior stabilized tOlal /8(9). 905-91~. condylar designs. CORR, 278. 147-132. Maloney, \\V.J\" SychlCrz. C., Bragdon, C. et al. (1996). Th\" Rosl,.·nhl·rg. A.G .. Andri.u:chi. T.P.. B\"In.kn. R.. d ..d. (1988 Olto Aufn\\l1c Award. Skeletal rcsponsl.' 10 well fixed P;ltdbr component bihlrL' in l\"ellll,.·ntl ....ss lot:d kne fel1lornl components inserted with and without cement. anhr{lpla~IY. Clill Onlt0fJ. 230. 106. C!ill Orr/lOp, Dec(333). 15-26. Manell. J.M .• Andriacchi. T.P., Rosenberg. A.G .. et nl. (1990). Rytldl. \\:.\\\\'. (19(\\61. F(ln:L'~ :Kling (In the fL'\\llVr:1! h.. a The rdationship b~t\\\\'een changes in pa!dl<lr heighl and prosthesis::\\ Sltltl~· CHI ~Il':lin g;lUg.e supplkd prostht.·st.·s function following total knee repl<lccment. Transcript of the 36rh AIlI1IUlI Mecril/!;, of the OrtJIO!Jlwdic Rt'search Soci- li\\·ing. pnsons. An(/ Orrhol' SCi/1Il1 (SIrf'f,[J. S~. 37, ely, 15(1).169. Morrison. J.B. (l9iO). The mechanics of the knee joint in rc- Schipplein. 0.0, ~ Andri:li.'I.·hi. T. (1991), Irlln:lcllOJl bClWl'C lation to nOfmal walking. 1 Biolllt·dlilllics, 3. 51-61. Murray, i\\'I.P.. Brewer, 8.1., &: Zuege. R.C. (1972). Kinesiologic activl' ~lIHI p:lssi\\'c kncl' SI:lhiliJ,er:-i during !L-\\'l' I walking. measurements of functional performance before and after Ort/lOp Res, 9. 113-119. McKec-Farrnr total hip rcpbcemelll. 1 BOlle 10ill1 Sllr~, 54;\\(2). 23i-256. Sci reg. A. L'\\: Ap,·ibr. R.J. (1973). Thl' pr,.-diction of Illuscul Noble. P.C. (1995). The design of cCfllcnlless femoral proslh~ scs. In J.J. Gallaghan, D.:\\. Dennis, & A.G. Rosrnb(,'rg lo:uJ sharing ~Ind joint fol'!':l,.·s in the 10\\\\'l'r l,.-'xtl\"l,.·lllilies du (Eds.). Ol'1hopaedic Kllo\\l'led~t! U,ulate: flip ((1Il1 Kllee Re:- cOllstructioll (pp. 12i-138). Rosemont. IL: AAOS. ing: walking. J Bioll/f.:dlCllliL~. 8. $9-102. Noble, P.c., Helmke, H.W., & Paul, J.P. (1993). Femoral posi- tion crilic;:t1ly affects the intl.'l'pretation of intra-vital hip Simon. S.R., Trit.'shmallll, H.W.. Bunit'll. R.G .. l·I.:\"I1. (198.3 force dnta. TrailS ORS. 18(2),443. Paul. J.P. (1976). Approaches to design: Force ':Ictions Irans- QU:ll1titati\\'e g:..lil allalysis ariel' 10lal klll..·..· i.lrthropb~t~- f mitted by join Is in thc human body. Proc I?es Soc LOlldol1. fllOl\\anicular dl'gl,.'lli.'rati\\·e arthritis . .I H(lI/~' loilft SlIrg. 6 192, 163-72. 605. Phillips, T.W\" Nguyen. L.T.. & Mull1'o, S.D. (\\991). Loosening Stallfrl,.'r, R.N .. Smidt. G.L. \\Vatl:-iwonh ..1.13. (1974). Clinic of cemen(\\css femoral stems: A biomcchanical nnalysis of 'l1H.I bioJl1l,.·c!wnical an'd.\\'sb of g;\\il I'ollowing Clwrnle.\\· 1 tal hip ri..'p!;U;l,.·Illl,.'1l1. CaRR. 99, 70-ii. St<'Tn. 5.1-1. &. Ins.a11, J.H. (1992). Postl..'rior :-Iahilizl,.·d proslh sis n.-sults \"ft,-r follow·up of nine to l\\\\·t.·!n' .\\\"l,.·;II'S. 1 BO 1oill/ SlIr!::, 7.,l:1. %0. Swany. .\\1.R. & 5\";011. R.D. (1993). Posll..'rjor poIYl,.·thylc \\\\,('ar in postnior cruci:lll..' lig;lllk'nt rl'laining !Owl kn an hrop1ast~'. J ..Inhrop!flsly. 8, ·B9~.J.J6. Taylor, 5.1., Walkl.:r, P.S., LlIlnon, P.L 1.'1 a!' (1997), Th orror<:l.:s in the dislal 1'l,.'Ill11r 'Illd knee during diffel\\.'111 acti itil,.·s llll':!sur\"d by klemetr,\\·' Tmllsl/clioJlS lilt' ·UI\"(I A lIt/al OUS .\\lc?l'til1::'. 259. Walker, 1).5. & Blunn, G.\\\\'. (1997). Ke.\\\"IIOIC Lel,.\"hu·L' II: !\\-\\o l'rll elL'sign of total kn('L' rt.'pl~ICl.:nll:lI!. In S. Niw::l, orYoshino. ~l. KUfOS;lb. K. Shino. S. Y;lln:l/11oto (l~ds.). R (,.·OllSffllctiOJ/ lite KlIt't' .1oil1t (pp. 129-1.J2). Toky Spri Il!!l'r- Vt.'rl:l!.!. \\\\\"aside\\\\~ski, R.C.~ Galante, J.O .. Lcii,dlly, R.\\I., cl al. (1994 \\\\'(::11' patll'rns on rt.'trie\\·\"d pnlyclh.vkllC tibial ill.~CnS an lheir relationship 10 Icchnical col1sidt.'l\"~lIiOllS during lCJI knt.'l,.' anhropl'lsty. CORR. 299. 31-43. Wcin:\\I1s. 1-1., I-Iuiskes. R.. .$: Groo!cnboer. H.J. (1992). Effec of mate!\"ial prol1L'l\"lies of femoral hip cornpOlll'nlS on bo rClllodeling. J Or/hop Rt's. 10. 8-15-53.
\\V~instdn. J.N .. Andriacchi, T.P.. &. Galante. J.O. (1986). Fac- Wilson. S.A .. McCann. P.D .. Gollin. R.S .. et a1. (1996). tors influencing walking and stairclirnbing following uni- prcherlsi\\\"(.' gait analysis in posterior-stabilized anhroplasty. J ..\\rthroplasty. 11{41.359-367. compartmental knee arthroplasty. J Arthropilisty. 1(2), Wimmer. ;\\1.A. &.: Andriacchi. T.P. (1997), Tracth'e force 109~115. ing rolling motion of the knt.'e: Implications for wear \\Vhite, S.C., Yack, 1-1.1 .. &. Lesswing, A.L. (1992). Pre- and post tal knee replacement..1 Biol/lcclulllics. 30(21.131~37. surgical gait of a total hip replacement p<:llielli. hilliS 71h \\Vright, T.i\\'1. &. Banel. D.L. (1986). The problem of s iI/lllual East Coast Gail ('olll('I\"('II('c 1992; Session 11, Rich- damage in polycthylL'ne total knee components. Cl {hop, 205, 67-74. moml Virginia. )'oder, S.A., Brand, R.A .. Pedersen, D.R., et a!. (1988). \\Villert. H. &. Buchhorn. G.H. (1993). Particle disease due to hip acetabular component position affects comp wear of ultrahigh molecular weight polyethylene. Findings loosening rates. CORR, 228, 79-87. from retrieval studies. Wear of polyethylene in total joint arthroplasty. In B.F. ivlorrey (Ed), Biological ,lIa/erial alld ,1Ice/wl/ical COllsideratiolls of Joillt Rcplacelllclit (pp. 87-102). New York: Raven Press Ltd.
Engineering Approaches to Standing, Sitting, and lying Chris 1. Snijders Biomechanics of Standing Reaching Biomechanics of the Pelvis Flat Versus Ball and Socket Joint Sitting Arm Rests Back Rest Seat Chair and Table Leg-Crossing Lying Sitting in Bed Decubitus Ulcers Summary References
Biomechanics of Standing Muscles extending from the neck to the ankle are A, lllu~tratjon of poor po~ture. B, lIlu~tration of go continually active to prevent collapse of the skele- toll. In view or this function, these muscles are po~ture. known as postural muscles. ObvioLisly, these I1lLlS· des are comparatively strong and do not tire easily. the neck. This manifes,s itself in postural de In skeletal musculature, postural and phasic mus- ration with the head placed forward, the sho cles-also known as slow and fast muscles-are blades raised and abducted, thc upper pan present (see Chapter 6). Thc mO\"phological differ- cel\"vical spine overextended, and the sho ences arc particularly marked in the Illuscles of moved forward. Figure 17-18 illustrates the birds, but in humans this distinction is less evident of strengthening of phasic muscles and stre because of the mixed composition of muscles. It is of postural muscles allowing for posterior assumed that the force proportion of postural to the pelvis, backward movement of the shou phasic muscles is 3:2. Exercise, especially participa- and elevation of the head by moving the chin tion in high-level sports or athletic training, can direction of the ears. change this proportion to even 5; I. Selective training of postural (red) muscles is ach-antageous because The unconstrained standing posture is these muscles are stronger and deliver more work studied because of the typical individual char than do phasic (white) muscles. The neurologist and istics. One of these characteristics is the loca muscle physiologist Vladimir Janda (1980) pointed the mass center of gravity (Fig. l7-2A, Ff=,F to the following relationship between postural and point of application of the resultant force o phasic muscles in agonist and nntagonist action: feet (FJ can be measured with thc aid of a plate (Fig. 17-28) and is called the center o I. Postural muscles tcnd to spontaneous func- sUl'e··(COP). A triangular plate with force tran tional 01' even anatomical shortening. This ers under the three corners can measure the manifests itself in a higher muscle tonus. tion of the COP as shown in Figure 17-2 D (Moll van Charante, Snijders, & Muldet; 2. Postural muscles have an inhibiting effect on Snijders & Verduin, 1973). The amounl of their phasic partners (inhibition of an antago- around a mean equilibdunl position reflec nist muscle by the agonist). 3. \\,Vilh insufficient variety in muscle lise, pos- tural muscles can be activalCd disproportion- ately, which leads to inhibition and weaken- ing of phasic muscles (pseudoparesis). This process can lead to an imbalance between groups of muscles, resuhing in poor posture with loss of mobility and increase of joint load. Figure 17-IA illustrates the distal crossed syn- drome according 10 Janda (1980). which involves shortened back extensor and hip nexor muscles versus weakened abdominal and buttock muscles. Anterior pelvic tilt and hyperlordosis are the result, which also manifests in walking with insufficient hip extension (normal is 5 to 10\"). A similar imbal- ance is the crossed proximal syndrome with short- ened pecton\\lis 111ajOl~ cervical trapezius. levator scapulae, and, less pronounced, the sternocleido· nlastoid muscles versus the weakened fixator mus- cles of the scapulae (rhomboidcus, distal and me- dial parts of the trapezius, and the serratus anterior muscles) and the deep llexor muscles of
COP nation of the smaller foot is greater than that of th D_ O_ Y larger foot. In practice. however, it is often th smallest \\\\'omen who wear the highest heels. In Fic IX Ix ure 17-4, the angle is 40°. Recom~1ended angles a Fg 10 to 14\", which translates to heel heights of 3 to y 0 cm maximum. F, In Figure 17-511, a chock is placed under a plaste cast around the lower leg and foot. This position o A BC the chock causes discomfort because it is placed i front of the ankle joint and causes a counterclock A, The resultant force between ground and feet {FJ equals body weight (Fy). B. Fv position (center of pressure path, wise moment, M = FI,.a. The foot and cast can be r COP) can be measured with a force plate with force trans- ducers under the corners depicting in a three-dimensional garded as a single entity, so equilibrium is obtaine force system (x, y, z). The person recorded in C rests more with a clock\\vise moment from force Fh exerted b on the left leg but has less sway than the person in D in x the tibia on the cast. A moment about the ankl and y directions. joint in the opposite direction occurs when the foo hits the ground with force F at heel strike (Fig 17-5). The heel of a shoe produces a lever arm wit respect to the ankle joint (Fig. 17-5(') that is consid Fs / •I I I I condition of neuromuscular control. Consumption A B of alcohol or certain 11lcelicines causes greater swaying. Marksrnen, conversely, stand significantly 0.24 II more stable. Most force plates have a rectangular design with four force transducers uncleI' the cor- I ners. Such construction, however, has the disad- vantage of static overdetermination. Because stand- F, I ing is not a pure static equilibrium as a result of I accelerations and decelerations of body mass dur- I ing swaying, the COP does not represent the exact + • position of the mass center of gravity' of the body. Recordings show that the COP is always located in 0.42 ( front of the ankle axis at the site of the as navicu- lads. Therefore, in standing, calf muscles are al- 0.18 I 0.72 I Fill ways active (Fig. 17-3, A & B). Figure 17-3, C & D 0 presents the location of forces in relation to foot C geometry taken from different literature sources (Snijders, 199\\). A, The foot-ground reaction force {F...} is always in front o the ankle axis (black circle). B, Therefore, the calf muscles Standing in shoes in general means standing on {FJ have to be constantly active. C, Mean values of lever uneven ground because a heel of elastic rubber has arms for calf muscles and foot ground reaction forces in a certain height for shock absorption and because at relation to foot length (t). The data is derived from diffe the ball of the foot, the sale must be thin to give ent sources. D, Lever arms for resultant forces on the hee goodllexibility. It is beller to look at the inclination (Fr.) and the forefoot (Fin) in relation to foot length (I). of the foot, not at the height of the heel, because the inclination determines the degree to which a person loads the forefoot. With an)' given heel height, the inclination of the foot depends on the length of the foot. vVith a certain heel height, the angle of incli-
AB c A, Heel strike produces a moment on the foot, which cl on the ground. B, A heel on a shoe increases this mome of the foot and the movement of the ankle. C, Cutting the back edge of the heel reduces the downward mom on the foot and produces a more horizontal movement the ankle. (F = ground reaction force at heel strike; a, a a\" = leverarm force F.) Diagram of a foot in a high-heeled shoe with a \"plateau sole.\" The angle of inclination of the foot is 40°, • erably larger than with a bare foot (Fig, 17-513), This A D moment claps the forefoot down on the ground, which can be reduced by! cutting off the back edge - ~B ...---.... of the heel (Fig. 17-5D). The laller produces a more horizontal movement of the ankle. This provision is 0)+ \\ ~ relevant for normal, sport, and orthopaedic shoes. C Excessive I110bility' or play is characteristic of an unstable joint, as compared with that of the healthy With three points of support in the top view (A) both in joint. In the following, instability refers to a Sllscep- the lateral (C) and the frontal view (0) of the foot, two tibility to tilting or falling as a result of interFering points of support are found at a distance ~rom each oth forces. In the top view of the foot (Fig. 17-6), one Such stable situations are not found in the dorsal (B) vi can identify three points where great forces are pos- with only one point of support at the heel. sible. Hence, both in the lateral and frontal view of the foot, two points of support are found at a dis- tance from one another. In the posterior vic\\v, how- ever, only one point of support at the heel exists. In this unstable situation, the heel must get stability' From the ball if the ankle fails, A stable positioning of the foot in a shoe means little if the shoe itself is unstable in relation to the ground because or a stiletto heel or a curved sole.
9 :e Free-Body Diagram in Stooped Postures The m,ISS cenler of 9t<1-:ll, 0: the 101<1l body hes '.i('nIGIlIy .1bo'J(' =the feCI Some lOlOl ,w:·s i)rc IndlCi)!e(! on lile skin. Z center oi !J\"wily: F, := upper body \\'if'IC);H lorce. f _-= body Inas~ r(,clCllon lorce; F. 0: bi:d: nlu:\".cli:· to:C(·. F.,;:: f(',lellon !OfCE' of the wel9ht =f()ru~; F. kilCiOI1 iorw (!f lhr bid', Inusc!e force: il = h:!'Jcr<1rlTl of upper budy '/,.'eiqht; b ::: 1('',',:r.'l(l'n of !Jdd: tT~lJ:;(I(~ force. A Calculation Box Figure 17-1-1. B Case A. The large r,)\"O b~twe('11 lever arms a onc: h (lllS:3'$ A, The arch of the foot compared with a Roman arch. I;;rgc· fnu .. c1(· :Q1C('''. m lhe IO'.'l('f i),Kk A<, an appfOXIIIl,1tI0n. F\" IS B. Flat joint surfaces in an arch experience compres- sion while shear is avoided. sao 0:N in a pr-r50n ~':ith Cl bodi h(:l!Jhl 130 ern and (1 bod.,. '.-/eI9I1, of 770 N \\;V:lh ii 27 en') \"nd b Sun, lhe U(lCk rmls- (Ie ior«~ (\"11 ik (iiku!ilt€.'d: F ivhf FilS >: 500 ,-, 2700 N . Body weight is carried by the arch of the fOOl, Calculation Box which resembles the form of a Roman arch (Fig. Figure 17-1-2. 17-7A). The arch mechanism relies on the firm con- nection between the two ends. In the root, this func- Case B. In the fH'X; C.15[o. the- le'.if'f ,,1m a is 11 COl• ...\"h,(;1 fe- tion is ascribed to the plantar aponeurosis. Schematic Sll1tS lf1 b,lCk musclE- force' F.. ~., alb. F.: - 11/5;, 500 \"., 1100 N. free-body diagrams of parts of the foot (Fig. 17-78) ThiS reduction of 60°/0 oi muSC,IIM forc(' illustrates lhClt keeping show the interesting feature of an arch: the joint re- lh(! Hunk more uprighl has c!rMrl,11ic mfluence on (still Ie) work action forces in the hind foot arc perpendicular to IO~l(1. wilh the le'/d iHn1 11 ilS 1110 1110s1 irnponarll iilCtor. Thus, a the flat joint surfaces. Thus, shea,· loading that mOfC' uprighl pOSlure rcdtJn~s ll1usclo and disc load consicleraoly. would tend to dislocate predominantly flat joints is avoided (Snijders et aI., 1993). Calculation Box Figure 17-1-3. Stooped postures often occur because of the role of the eyes in combination with the work of the L hands. This raises large muscle forces in the lower back as a result of the large ratio of the lever arm (a) of upper body weight (F,) and the lever mm (b) of back muscle force (F). A cross-section at the level of L5-S 1 shows that the mass center of gravity of the body part above this cross-section is situated at the armpits (Calculation Box 17-1, Case A). The equilib- rium of forces in the verlical direction shows that the weight force (F~) equals its reaction Corce in the spine (FI1:). It is assumed that this reaction force acts in the middle of the intervertebral disc (the axis of .
a optimal because the)' are perpendicular to th l·tQ longitudinal orienlation of the spine. Howeve the sacroiliac (51) joint surfaces are parallel to th ~~ largest forces and are not protected against disl //11111111 1111/11111 cation bv the closed form of a ball and sock joint. Thus. the 51 joints arc vulnerable to shea ing because of their predominant flat surface which arc almost parallel to the plane of maxim load (Fig. 17-9), In a biomechanical model. Sn .iders et al. (1993) made the assumptiollthat in < dition to strong ligament support, muscle forc B are always required for compression of the S A, An obstacle prevents the feet from advancing any fur- M ther, leading to a forced posture. B, Same reach distance, but a more comfortable position, thanks to suitable support ~ of the pelvis. The lever arm a is reduced by 30%. and the F arms and eyes are in a better position. (F\" = center of gravity Scherr,atic drawing illustrating that the sacroiliac joint is upper body; a = lever arm center of rotation spine to F\".) vulnerable to shear when loaded in the sagittal plane (m ment M/Force F). Integrity of the joint reli~s on ligament • support; joint compression by muscle force can raise fric- tion and prevent shear. this joint), F~ and F,,! form a clockwise rnomcnt with ':; M = F(:.a. The equilibrium of moments comes from the couple formed by back muscle force F, and its reaction force F\" on the disc. The total force on the disc is the (vectorial) sum of F\" and Fo.' The equi- librium of moments yields: Fwa = F\".b, So. the total muscle force Fs = alb, F,:, \\ REACHING Conditions like those shown in Figure 17-81\\ occur at work when the person reaches as far as possible to a distant object while the feet cannot be placed any further forward, either because of the design of the machine involved or because the tocs touch a tank with liquid. The provision of a support for the pelvis as shown in Figure 17-88 allows for the same I'cach but with greatly improved posture. This mea~ sure reduces the lever arm by approximately 300/0 and the arms and eyes are in a much more suitable position. In the workpl<lce, people often seek sup- port 1'01' the pelvis or upper legs by using the edge of a tablc or bench. BIOMECHANICS OF THE PELVIS With reference to Figure 17-8, the largest forces always act in the longitudinal direction of the spine. The position of the intervertebral discs is \\1 i%:, ':' ~\".~ ~ U'.~ .... ly,=~\"\"·,z':liC,::c:::c_*i?\"~;::se
JOInts to raise friction, which resists shearing. 2 ---I--\\--\\ II J/II -1---5 This so~called self-bracing is also expected in sit- ting with reference to ligament creep by the load 3 IIIIII 6 of upper body weight. ~IIIW FLAT VERSUS BALL AND SOCKET JOINT 4--+..,..,.~ The different joints in the human body have a wide variety of often highly irregular forms. By compar- Transversely oriented muscle forces that cross the sacroiliac ing a ball and socket joint (or cylindrical joint) with completely flat joint surfaces, some main (51) joints can press the sacrum in between the hip bones. principles of load transfer can be identified. If a transverse force is applied to the bones in Figure Such forces caudal to the SI joints may contribute to the 17-1011, shearing must occur to produce stress in ligaments, which stops the movement. The bones mechanism of the pelvic arch by opposing lateral move- are no longer in line. A ball and socket joint (Fig. 17-10B) is better protected against transverse ment of the hip bones. (1) linea alba, (2) external oblique forces. \\\"'hen a flat joint is loaded with a bending moment (Fig. 17-IOC), the joint reaction force can abdominal, (3) transverse abdominal, (4) piriformis, (5) rec- shift to the edge of the joint surface and can pro- duce with ligament or muscle force F{ a couple I tus abdominis, (6) internal oblique abdominal, (7) ilioin- ~~~_~~_ with moment M = Ff .2c \\Vhen a ball and socket 1.-_9_u_i_na_I_I_i9_a_m_e_n_t,_a_n_d_l_8_)S_I_iO_i_n_t. joint is loaded with a moment (Fig. 17·10D), the joint reaction force cannot shift to the edge of the joint surface. Here, F\" and F1 form a couple with moment M = Fi.r. This shows that a flat joint can transfer a \"factor two\" larger bending moment than a spherical or cylindrical joint at the hip joints. In Figure 17-10, C d.! D, the joint reaction force is perpendicular to the tangent to the joint surfaces at the bone contact point. This holds when friction can be neglected. At the spherical hip joint the problem of a lesser moment of force has n-··-u '-H li:~F' ~~lF, been removed by' means of the trochanter rnajoc Il-· ,I. 2r ,I ~. r .I which enlarges the lever arm of Illuscles consider- abl.\\'. From this comparison, it can be concluded A B C0 that SI joints are vulnerable to shear but can trans- fer large bending moments and compression (Sni- A, Forces near and in the plane of a flat joint result in jders el aI., 1993). shearing before this movement is stopped by ligaments. B, A ball and socket joint is well protected against shear. rVlusc1e forces that cross'the 51 joint surfaces C, Because of the greater lever arm, a pure bending mo- can produce compression, in combination with ment can be better transferred by a flat joint than by a Forces in ligaments and fascia. This protective sys- ball and socket joint (0) (F., delineates joint reaction tem requires the concerted action of muscles in force, Fi is the ligament or muscle force, r represents the the back, pelvis, and legs. Some muscles with the lever arms). appropriate transverse direction arc indicated in Figure 17-11. SI joint stabilit:, can also be ascribed to the 111cchanism of the pelvic arch, which relies on the protection against lateral movement of the hip bones. Figure 17-12 compares the arch of the fOOL and the arch of the pelvis. The effect of a pelvic belt distal \\0 the 51 joints also can be seen (Snijders et aI., 1993). The pelvic arch receives bi·
Analogy between the arch of the pelvis and the arch of arm rests. the arms arc placed on the table, cross the foot. This shows analogy between the load on the before [he chest. or laid in the lap. 1n cars, arm re sacroiliac joint and the load on a tarsal joint. The (horizon- arc ohen absent or lOa low. Drivers tend to turn tal) line of action of a pelvic belt distal to the sacroiliac the side of the door and put their arm on the edge joints is also seen. Reprinted with permi.ssion (rom Snijders, the window. The other arm rests on the lap or is p C.J.. Vleeming. A., & Stoeckarr. R. (1993). Transfer of lum- on the back rest of the passenger seat. Because of t weight of the arms, the driving wheel is often held j bosacral load to iliac bones iJnd legs. Part I: Biomechanics of the lower side. A proper arm rest must be placed b self-bracing of the sacroiliac joints and irs significance for ((eat- low the mass centers of upper and lowcr arm. A su men( and exercise. Clin Biomechanics. 8. 285-294. pan at the wriSl is lherefore useless. In cockpit sea arm rests arc adjustable in height but only by tlu-ni lateral support from the hip joints. When one sits about an axis adjusted to the back rest. This co on the ischial tuberosities, the supponive forces stnlction is a logical error because in a high positi are below the SI joints and the mechanism of the the arm rest only supports the wrist and leaves air b arch is absent. low the elbow of the tall pilot. Sitting Postural Problems at VDT Workstations It is possible to predict whether a chair is comfort- Afemale computer operator works at a VDT worksta- able or not. independent of the size of the pel·son. Evaluation of design criteria on the basis of biome- .... tion that includes a corner CO!1iputer desk fixed at a chanical aspects demonstrates the poor design of certain height, an office chair without adjustable features, much furniture. Body posture is highly innuenced CPU. monitor. keyboard, and mouse. The VDT operator by the form of a chair. A good sitting posture is complains about pain in the right shoulder, wrist, and low characterized by minimal muscle effort. which is back region. The work setup necessitates excessive arm produced with proper support by arm rests, back resl, seat, and fOOL rest. Nonetheless, the first law of extension that puts an increased load on the shoulder good scating is the ability to change posture regu- larly (Case Study 17-1). , joint and mechanical pre.ssure on the elbow and forearm, affecting underlying soft tissues of the arm. The unsup- ported sitting posture shown in the figure below creates inc_reased disc pressure, particularly in the lumbar spine. ARM RESTS Case Study Figure 17-1·1. The importance of arm rests is often underestimated. Arm rests unload the shoulder girdle, which is a loose construction hanging on the spine by means of liga..: ments and muscles. The weight of the arms is 100/0 of body weight, which is considerable. In chairs without
BACK REST tion. (\":'urther transhltion results in backward tilt o The back rest provides stability for the vertically the pelvis and lumbar k~·phosis. The reader can cas erected trunk, analogous to Figure 17-8B. HoweveI~ il~' perform this test. The fUllcticl!l of a lumbar sup in prolonged sitting, the prevention of a lumbar port is to exert a firm force on the upper side of the kyphosis seems to be the most important function of pelvis ..md the lumbar area to prevent tilting of the the back rest. The significance of a lumbar support spine into kyphosis. This support should not reach can be illustrated with the so-called click-clack phe- higher than the lower edge of the scapulae. The tho nomenon (Fig. 17-13) (Snijders, 1970, On the form racic spine is still enough (ribs) and a higher lnlC of the human thorneD-lumbar spine and some as- rest pushes the shoulder blades rorward, which pects of its mechanical behavioc Thesis, Eindhover). \"overrules\" the lumbar support and hinders th This can be experienced when one sits upright on shoulders to stretch and to turn to the lert and th the edge of a straight chait: Slow forward translation right. The absence of a back rest, like sitting on of the trunk increases lumbar lordosis, whereas slow crutch. alwa~'s leads to a C-form of the spine. backward translation moves the center of gravity above the ischial tuberosities into an unstable posi- SEAT oo The weight of the trunk, head, and (part 00 th arms is alnH)st completcl~' carried b~' the ischia The lumbopelvic click-clack phenomenon. Lumbar lordosis tuberosities. This leads to a simple representation o is the result of a mass center of gravity of the trunk at the the static equilibriurn of the upper bod\\' in sitting ventral side of the ischial tuberosities (right). The mass cen- The aim of Figure 17-14 is to illustrate that a hori ter of gravity of the trunk dorsal to the ischial-tuberosities zontal seat always raises friction at the ischia forces the lumbar spine into kyphosis (left). The latter can tuberosities, tl1<.\\t friction can be completely elimi be prevented by the exertion of lumbar support on the nated b~' means of a moderate seat angle. and tha upper side of the pelvis and the lumbar spine. the angle betwL'en seat and back rest is ()ptimal be twcen 90 and 95°. which is met b~' the classical rock ing chair (Snijders. 1988). In Figure 17-1413, a free bodv diagrarn of the trunk is given, with cross-section at the upper legs and SUppCl!'t forces a the seat and the back rest. The mass center of grav it}· of this part of the body is near the ~lnnpits. Th line of action or the gravil<.ltional f()rce (FJ and th back rest (Fb) intersects at point S. The resultan force exerted by the scat on the ischial tuberositie (FJ must also intersect at point S for static equilib rium. The magnitude of the three forces is f()llnd b the composition or the tri ..mgle of forces. It can b concluded from the direction of Ft that a back res force always raises a hOI'izontal force componen (Fh) at the ischial tuberosities. The consequence friction on the skin and )undcrl~'ing tissues tha causes discolnt\"ort. Fronl Figure 17-14C. it can be seen lhat with slightly inclined scat surface at the ischia tuberosities, the resultant seat force (FI ) stand nearly' perpendicular to the surface. There is a re duced force component in the shear direction \\Vhcn the back rest is tilted further) the seal angl must follow because FI becomes more inclined. [ Figure 17-15. the relation between back rest angl and scat angle is given for the condit.ion that Fric tion is absent (Goossens & Sni.iclers, 1995). This re lation was found with a biomcchanical model o
o Measured - - Calculated 20r---------------- 1ii '\"(f) o'--~......\"O:_~___,~~,__.,;;,____;;,___;;.,____::':::_ 70 72 74 76 78 80 82 84 86 Backresllnclination (degrees) Relation between back rest angle and seat angle. No fri tion exists at the ischial tuberosities. The solid line is fro a model calculation. The vertical bars represent measure ments (n --0 10), including tall and small persons. AB the total bod)-' divided in five links and \\vas verif by measurements on 10 healthy subjects with d ferent body height. Back rest inclination is lar in auditoriums, cars, and eas)-' chairs at home. T facilitates a horizontal direction of looking and h the advantage that the back rest contributes in c rying the trunk. Head rests cannot carry' the he when they are positioned too far fonvard. Fig 17¥16A illustrates that the head rest can only g F, Fgr( c D 1:, ~\"FIG:~17-14 ~ AB ~~--~~ A. The mass center of gravity (represented by a circle a gravitational force F~) is behind the atlanto-occipital jo A, Sitting with maximal support. B, Free-body diagram of axis A-{Ft represents the joint reaction force). With sup the trunk. Back rest force (Ft» and upper body weight port of a head rest, no muscle forces are needed for e force (F() intersect in the point S. The resultant seat force librium of the head (F'J = Fl ). B, The head rest is posi- at the ischial tuberosities (f\\) must intersect in 5 as well, tioned too far forward and is useless. Muscle force F, and therefore inclines backward. The horizontal compo- remains needed for equilibrium andF, increases. nent (Fh) must come from friction acting on skin and un- derlying tissues. (-0, By tilting the seat slightly backwards ((\\), the resultant seat force (fl ) becomes nearly perpendic- ular to the seat surface and friction on the ischial tuberosities is significantly reduced.
support when the mass center 01\" gravity of the Upright Position Anteflexion Position head is dorsal to the axis of the atlanto-occipital joint (A). In Figure 17-168, the head rcst hinders movement because the head cannot bc tilted back- ward. This design error can be recognized in trains \"nd buses in which people nod or rest with the head sidewa.ys. CHAIR AND TABLE Sagittal r' view A chair is important for a good posture, but when ~~I~ tasks such as reading and writing are involved, the hcight and inclination or thc desk or table playa ~Fj dominating role. Despite good chairs, postures with the back benl, sagging, or twisted can be ob- Transverse served. Problems arise when such extreme pos- view tures are mainrained for long periods of time. In the relaxed posture according to Figure 17-17;\\ the Anteflexion of the head leads to a larger lever arm of the direction of vision with the head upright is given weight force (F'l) with respect to the atlanta-occipital joint (dOlled line). However, reading with the trunk up- axis (A) than in the upright position. For equilibrium, right requires a sharp bend in the cervical spine, which cannot be sustained for a long time. Fur- greater neck muscle forces (FJ are needed. Forward indi- thermore, in this position the reading distance is larger than 25 to 35 CI11 (the height of the table nation also raises tension in the ligamentum transversum must coincide with approximately the level of the atlantis (F 1) to transmit the force from the dens (Fel), which elbows). Consequcntly, children and adults always is directed transversely to the cervical spine. This force fol- bend fonvard with a curved spine over the table lows from the equilibrium of the atlas, with joint reaction (Fig. 17-178) and urgings to sit up straight are forces F1 and Fl· useless. In the past, inclined desks wcre common. The fixed angle was approximately 120 because at a higher angle, paper will slip downward on paper be- ,,-- ,: A Bc ncath it (Snijdcrs ct al.. 1990). A steeper angle also is inappropriate for arm suppon. Antcnexion of the A. Reading while sitting upright demands, forced by the head leads to increased neck muscle rorc~ with al- eyes, a bend in the cervical spine. Moreover, the reading Ill()S~ a factor of three. Furthermore, ligaments such distance is too large. B. Therefore. the back is bent. C, A as the lrgamentum transversum atlantis (Fig. 17-18) proper height and a small inclination of the desk (up to are loaded considerably in anteOexion, while this 12°) brings the page to the eyes, instead of the reverse. load is almosl absent in the upright position (Sni- .iders et al.,1991). In a rorward-bent position or the trunk and head, considerable muscle forces arc needed. A pro-
( Pain 23 4 Tissue \\ breakdown Muscular Poor spasm t lL Smooth \\ blood supply 5 i' ,l::! 6 J ~•I • 7 Rough 8 The vicious circle of pain. • longed, continuous contraction (above approxi- 9 11 12 13 mately 20% of voluntary maximal contraction) ~~ causes a lack of oxygen, accumulation of sour metabolites, and intracellular shortage of potas- ~ sium. Pain is the result, which can lead to muscle spasm, which closes a vicious circle (Fig. 17-19). 14 15 Prolonged isometric contractions can even cause an 16 17 18 19 inflammation process with fibrosing; the result is a ~ 21 22 23 passive shortening of the respective muscle. Enter- ing this vicious circle is also ascribed to cold and 20 stress. Attempts to cut this circle are analgesia, mus- 24 cle relaxation, and improvement of circulation by Examples of poor chair design leading to poor posture. massage and heat radiation. These efrorts involve the treatment of symptoms, while the improvement of the work posture addresses the origin of the mus- cle spasm. Some biomechanical problems in sitting arc illustrated in Figure 17-20. I. The absence of an arm rest loads the loose shoulder girdle, which hangs from the spine with ligaments and muscles. 2. Sitting without a back rest always leads to a C-form of the spine. 3. A back rest must leave space for the but- tocks. \\ViLh a straight back rest staning from the seat, no proper support can be given to the upper side of the pelvis and the lumbar spine. A support that is too high does not provide force on the lumbar area and a C-fonn of the spine is the result. 4. If the back rest is too high (as it is in some folding chairs), it gives no support to the lumbar area and the back bends in a C-form.
5. A vertical, straight, and high back rest also molded lO the individual anatomy should be promotes lumbar kyphosis. restricted La the severely handicapped. 21. A head rest lh<:tt is placed too far forward 6. On a horizontal slippery scat, the botlom hinders upright sitting. slides forward and the body sags in the 22. Tall people experience hindrance from a chair because the [detion force (Fh in Fig. head rest thm pushes against the shoulder 17-14A) is absent. (analogous to 5). 23. \\,y'hen reading and writing at a horizontal 7. On a rough horizontal seat, the friction desk, adults and children always show force (F\" in Fig. 17-14B) causes discomforl; flexion of the thoracolumbar spine. This people avoid back support and adopt a pos~ forward bending is independenl of the tureasinFig.17-17B. chair. 24. Neck and shoulder complaints arc reponed 8, When silting on a horizontal cushion like on with visual display units. \\.Vhen it is not nec- a lounge chair, the cushion is moved by the essary Lo look at the kevboard all the time. horizontal force component at the ischial the optimal height of the screen is approxi- tuberosities (reaction force of Fh in Fig. mately al eye level (Snijders el aI., 1991; 17-14B) and mOves inlo Ihe room. Wall el aI., 1992). 9. Further lilting of Ihe back rest must be cou- LEG-CROSSING pled to the same increase of scat inclination. In ailvlancs, ho\\\\'evel~ seats remain almost The majority of people cross their legs often when horizontal. sitting, alternating left over right and righl over left (Fig. 17-21). There may be many' reasons to do this, 10. Broad persons are jammed by lateral sup- bUl in the literature th(~rc is no scientiflc proof pons. In common cars, such pronounced aboLlllhe benefit or demerit of leg-crossing. One e:\\'- lateral supports arc useless because curves are not taken with more than 30 111/S1 (0.3 g AB c centripetal acceleration). A back rcst with a uniform moderate curvature flts everybody Even in sitting with the support of back rest and arm rests, and provides sufficient lateral support. the internal oblique abdominal muscles are significantly more active as compared with when supine. Crossing the 1I. Elderly people do nol appreciale a low or legs (upper legs crossed or ankle on knee) lowers this activ- deep sinking seat. ity. Therefore, leg·crossing is assumed to be functional and should be allowed for in work places (except car seats and 12. The length of the seat in a chair for adults is cockpit seats). too large for children. 13. The lengths of scats in easy chairs or lounge chairs arc often too large for adults. 14. Although the length of Ihis seal (2) seems to be sufficient. Ihe effective seat length (t) is fm\" too small because of a useless elevation at the dorsal side and a too large radius at the foreside. A radius of approximately 30 to 50 mm is sufficient. 15. Pronounced lateral supports on seats have no use, also common in cars. Broad people experience painful jamming on the thighs. 16. A back support that is too low cannot give sta· bilily 10 the LJunk (analogous to Fig. 17-88). 17. Arm rests are even more important in easy chairs than in straight chairs because the arms cannot rest on a table. 18. Easy chairs exist with a pronounced C-form. 19. A seat Ihal is 100 high causes pressure on the thighs. whlch promotes \"sleeping legs.\" 20. Seats must be straight in the anteroposterior direction. The curvature in the lateral direc- tion should remain moderate. Forms
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