limb

Lower Limb Prostheses: Design Considerations



John Register’s Olympic dreams were crushed while he was training for Atlanta



in 1994. An improper landing from a hurdle jump hyper extended his knee, severing a



leg artery and ultimately leading to an amputation. Register was faced with a choice



between confinement to a wheelchair or a prosthesis followed by intense therapy and



rehabilitation. The desire to be independent and mobile led him to opt for the latter.



Register was fitted with an above-knee prosthetic leg. The limb consisted of soft,



flexible plastic and carbon graphite with openings to allow the thigh muscles to grow.



This was revolutionary compared to ordinary prosthetics, with their hard, rigid sockets.



Non-pliable materials confine the muscles, causing them to atrophy. Prosthetic



composition was a necessary consideration for John Register; compliant substances



absorb the impact of heel strike, and the muscles can generate the necessary energy to



move the knee and proceed through the stride. If the material is too compliant, there is



not enough reaction force to propel the leg forward into swing phase. In this situation,



the knee and hip must generate a large amount of force to continue the motion, costing



the amputee a great amount of extra energy and work.



About 344,000 people in the United States have lower limb amputations, and



approximately 116,100 are added to this population each year. Causes include disease,



trauma, birth defects, and tumors [1]. These statistics demonstrate the importance of



prosthesis and the need for continuous improvement for comfort and ease of use. Much



research has been conducted on the mechanics of the leg, and manufacturers have



designed artificial limbs to mimic the human leg as closely as possible. However, further



development is always within our reach.









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This paper looks at factors affecting how and why prosthetics have unnatural gait



and how research is trying to correct for these deviations. Factors discussed are the basic



structure of a lower-limb prostheses, materials, weight and mass considerations, power



requirements, biomechanics, and tradeoffs in motion and stability. By understanding the



considerations taken into account when designing prostheses, we can better realize the



complexity and importance of prosthetic design.







General Structure of Lower-Limb Prosthesis



The standard prosthetic design utilizes the interaction between a tibial spine and



femoral cam. The spine-cam contact substitutes for the posterior cruciate ligament (PCL)



found at the back of the knee. The primary function of the PCL is to prevent the tibia



from slipping backwards on the femur and to allow for some resistive force between the



two. The spine-cam interaction causes the femur to roll posteriorly on the tibia (called



femoral rollback) to increase the amount of knee flexion. The main difference between



designs and manufacturing specifications for various lower limb prostheses lies in the



unique geometries of the spine and the position of the cam relative to the tibial spine.





Materials



The typical prosthesis is made of a metal alloy and high-density polyethylene.



Other materials used in knee prostheses include carbon fibers, aluminum, titanium, and



foam. Polyethylene is widely used because it is able to withstand continual forces



without significant wear that would require it to be replaced frequently. The degree of



wear is dependent on the amount of motion, the quality of material, and the roughness of



the tibial base plate. If the prosthesis is extremely rigid and limits the amputee’s range of







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motion, the polyethylene wears down faster. This is important for manufacturers to bear



in mind when considering a specific prosthesis. A restrictive design is safer due to its



increased stability, but material wear occurs faster than in a more lenient joint. Injury



and fracture can occur with extreme stiffness.



Most leg prostheses have an upright tibial spine and an oval femoral cam to



mimic the knee joint where the femur and the tibia would normally interact. A net



downward, compressive force is put on the tibial spine to produce an impulsive reaction



force and push the leg forward. Without any support, the loading shock would not be



absorbed. Material would wear quickly and the amputee would have unnatural gait.



However, too much absorption (a soft limb) would delay the stride due to insufficient



reaction force upon loading. To support the loading force while providing a sturdy



surface to create an impulsive force, bone cement is often used for its durability and



stability characteristics.







Weight and Mass Considerations



The optimal weight of prostheses components has been disputed for many years



and depends on the materials used, the type of prosthesis, and the requirements of the



user. Some prostheses weigh less than 2 kilograms [8], which some may argue is too



light. A lighter prosthesis has the advantage of requiring less energy to move, therefore



involving less muscle cost. However, unequal limb weights (between the normal leg and



the prosthesis) can lead to gait asymmetry due to different centers of mass and mass



moments of inertia. Additionally, a lighter prosthetic requires more control during swing



phase, which may offset the energy saved in generating motion. As the mass decreases,









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the inertial forces to stop it increase because the limb acts like a pendulum and does not



stop until a force opposes its motion. Thus, the lighter the limb, the harder it is to control



once in motion.



Some may contend that heavy prosthetics can be detrimental. As the limb



becomes heavier, more power must be produced at the hip to execute swing phase. A



heavy limb can be very tiresome for the user. With extreme weight, the prosthesis may



“drag” and slow down the amputee.



The positioning of the prosthesis’ mass should be taken into consideration to



determine the best conditions for the user. If most of the mass is near the knee, then the



effect of added mass is minimal. However, if mass is added distally (further from the



torso, usually at the ankle), then the effect is larger and disadvantageous for natural and



efficient gait. Research aimed at maximizing gait symmetry has shown that as prosthesis



mass increases proximally (nearer to the torso), stride time and single limb stance



increase significantly, indicating more stability [8]. This means that if more mass is



located at the knee rather than the ankle, the user will be steadier. One way to visualize



this phenomenon of mass placement is to imagine spinning with a ball on the end of a



short string versus a long string. The longer the string is, the farther the ball is from your



center of mass. As the object goes farther from your center of mass, the harder it is to



spin at the same speed and energy level. You can spin faster, use less energy, and



maintain better control with the ball on the shorter string. Therefore, the farther the



heaviest mass of the prosthesis is from the body’s center of mass, the harder it is to



control that mass and use it efficiently. A larger mass located at or near the socket of the



prosthesis is therefore advantageous for control and efficiency.









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It has not yet been possible to perfectly replicate normal human gait with



prosthetics. Technology and research minimize deviations while also considering



comfort and energy. Amputees may show altered characteristics of gait such as increased



swing and step time, increased step length, and decreased stance time solely on the



prosthesis [6]. The increases associated with stride are a result of different energy



requirements between the natural and amputated legs. More work must be generated at



the hip of the amputated limb because there is no assistance from the knee, ankle, or foot



to propel the stride. Less time is spent on the prosthesis due to instability. To keep one’s



balance and walk more naturally, the true limb occupies more time in single stance than



the prosthesis. This allows the user to make adjustments in balance as necessary.







Power Analysis of Gait



Walking requires a substantial amount of power to be generated at different



portions of the leg. The three main locations of power production and absorption are the



ankle, knee, and hip. Amputees lose power at their ankle because they lack toes and an



ankle for push off. Plantar flexor muscles in the foot and toes are used to propel us



forward, and we catch ourselves on the heel of the opposite leg. Because amputees lack



these muscles, they must produce the power elsewhere, usually to a greater degree.



The knee is a small factor in power analysis because the least amount of power is



generated and absorbed at the knee throughout the entire gait cycle. There is little



difference in knee activity during loading, extension, flexion, and swing phase when



comparing normal and amputee gait. Amputees are slightly disadvantaged because there









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is not a continuous flow of energy throughout the leg. A small amount of energy is lost



at the knee but is not significant enough to alter gait substantially [3].



The main source of difference in power production between normal and amputee



gait exists at the hip. The hip is generally a large site for power generation due to the



large surrounding muscles and its proximal location to the body’s center of mass. The



hip flexors must compensate for the absence of the feet and ankle muscles to “pull” the



leg in the desired direction. This can lead to wearisome and asymmetrical gait for



amputees unless prosthetic designs can conserve energy and increase gait efficiency.



Compare the amount of energy it takes to walk normally versus the effort required to lift



the leg, swing it forward, and land flat on your foot without any ankle flexion or push-off.



The knee is not able to assist in this action, whereas it would normally carry through the



stride. More mass at the knee rather than the ankle allows the knee to swing forward



easily to progress to the next step, yet is closer to the body’s center of mass for control.



The feet and ankle muscles control and utilize leg power efficiently.







Prosthetic Biomechanics



The keel is a prosthetic foot. It is crucial to the mimicry of normal gait because



the length of the keel determines the timing of heel rise. Heel rise is defined as when the



ball of the foot and the toes are still in contact with the ground, but the heel lifts to push-



off. The keel provides flexibility to the amputee by minimizing the stiffness of the



prosthesis. Oftentimes the prosthesis’ rigidity is the most criticized aspect of the design.



However, some stiffness is required for stability during stance. To balance safety and



comfort, the keel is designed to offer some flexibility by acting like a foot.









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The ground reaction force (GRF) is the force the ground imposes on the standing



leg when loading occurs. If the GRF passes in front of the keel, the heel rises earlier.



This occurs because more weight is put on the front of the foot. The heel must



compensate by lifting and allowing the foot to progress forward. The opposite is true as



well. If the GRF is located at the rear of the keel, the heel does not rise unless the weight



is shifted to the front, moving the GRF to the front. This is analogous to standing on the



balls of your feet versus on your heels. If your weight is mainly on your heels, heel rise



will not occur until you shift your weight forward.



If the GRF passes through the keel itself, the leg does not move. This situation



can be understood when standing in place. The GRF is centrally located, leading to a



balance of forces in the posterior and anterior sections of the foot. Therefore, there is no



movement unless the weight is shifted. This stabilizes the prosthesis and retards shank



advancement for more natural gait. If the limb were not restrained by the location of



forces, it would not allow the amputee enough time to stabilize herself.







Motion and Stability Tradeoffs



There is a tradeoff between maximum knee flexion (which provides the largest



range of motion) and prosthesis stability. Small increases in knee flexion can cause



significant decreases in stability. Without the PCL restraint to limit knee flexion,



dislocation would occur frequently.



Dislocation occurs when the femoral cam translocates forward over the tibial



spine. This dislocation is most likely to occur when the knee joint is put into a position



of forced flexion. Two examples of forced flexion include standing up from a low chair









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and putting on shoes while standing. Forced flexion allows the knee to pass beyond the



tibia while the tibia reacts by creating a force equal and opposite to the loading force.



The compliance of the ligaments and other soft tissues surrounding the knee also prevent



dislocation by absorbing some of the force. Without a balance of forces, stability is lost



and dislocation occurs. For the amputee, the absence of a true tibia and soft tissues



results in an acute dislocation in which the knee is locked in a flexed position. The



prosthetic knee is most likely to dislocate at the maximum flexion angle [2]. Therefore,



prosthetic designs must maintain stability but allow maximum knee flexion angle without



dislocation.



Changes in design are measured with the dislocation safety factor (DSF). This is



a geometric parameter representing the chance of dislocation. The DSF is the vertical



distance from the top of the tibial spine to the bottom of the femoral cam. As the DSF



increases, the femoral cam moves below the tibial spine head. This movement decreases



the chance of dislocation. Thus, the higher the DSF, the less likely the limb will



dislocate.



The DSF peaks at the middle of the knee flexion range and decreases with either



continued flexion or extension. Maximum knee flexion is found to be around 125



degrees, with the highest DSF at about 70 degrees of flexion [5]. Geometric alterations



of the prosthesis can affect the DSF. It is by this geometric variable that multitudes of



prosthetic designs have come about. The DSF increases with tibial spine height and also



when the spine is placed slightly forward. This positioning allows the spine and cam to



interact at greater flexion angles. However, forward positioning of the spine decreases



maximum knee flexion angle due to femoral rollback. Therefore, a balance must be









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found to optimize the DSF for stability and maximize knee flexion angle. Most amputees



find that 115 degrees of knee flexion is sufficient for conducting everyday activities [5].







Conclusion



Prosthetics maximize amputees’ independence and allow them carry out their



daily lives. For many like John Register, it is unfathomable to lose the ability to walk.



By designing prosthetic legs, immobility can be avoided for most amputees. With much



rehabilitation, amputees can walk with almost normal gait and perform activities like



those they performed with their true legs. Many factors must be considered to meet the



individual needs of each person, leading to multitudes of prosthetic designs. Natural and



efficient gait is amazingly difficult to replicate due to the complexity of the human legs.



Features such as materials, weight and mass, power, biomechanics, and motion and



stability must be accounted for. Thus, prosthetic research and design is a rapidly growing



area of biomedical engineering and serves to accomplish one main goal: to improve the



quality of life.









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Bibliography





1. Amputation. Handbook of Disabilities. University of Missouri and RCEP7.

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stability in posterior substituting knee arthoplasty design. Journal of

Biomechanics. 28 (10): 1155-1166.

3. http://moon.ouhsc.edu/dthompso/gait/epow/pow1.htm. 01/2003.

4. http://moon.ouhsc.edu/dthompso/gait/pobmk/base.htm. 01/2003.

5. Kocmond, JH, Delp, SL, Stern, SH. 1995. Stability and Range of Motion of

Insall-Burstein Condylar Prostheses. The Journal of Arthroplasty. 3(10): 383-

388.

6. Lin-Chan S, Nielsen DH, Yack HJ, Hsu MJ. Changes in gait performance at self-

selected walking velocity with proximally added prosthesis mass in persons with

transtibial amputation. University of Iowa, Iowa City, IA. 2001.

7. Scuderi, GR, Pagnano, MW. 2001. Review Article: The rationale for posterior

cruciate substituting total knee arthroplasty. Journal of Orthopaedic Surgery.

9(2): 81-88.

8. Selles, RW. Optimal weight of below-knee prosthesis. Institute of Rehabilitation,

Erasmus Medical Center, Rotterdam, the Netherlands.

9. www.americanplasticscouncil.com/athletic/sports/register.html. 01/2003.

10. www.maitrise-orthop.com/corpusmaitri/orthopaedic/88_afriat/afriatus.shtml.

01/2003.









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