-In 1912: The first surgical procedures to treat hip osteoarthritis were interposition grafts. Sir Robert Jones placed gold foil within the joint space, others used native soft tissue like fascia lata or synovial tissue.

-In 1923: Dr. Marius “Mads” Smith-Petersen pioneered the Mould-Arthroplasty procedure in Boston.  The technique placed a glass mold over the femoral head to stimulate fibrocartilage metaplasia and form a smooth articular surface (the idea came to him after discovering a smooth synovial membrane surrounding a glass shard pulled out of a patient’s back).  Not surprisingly, the glass broke after a few months, however, the very-short-term results were positive, and thus more durable materials were utilized (first pyrex, and then Vitallium…a metal alloy, see below).  [23].  It was the first attempt at a surgical procedure to restore the joint instead interposing tissue.

-In 1932: Dr. Charles Venable and Dr. Walter Stuck, after studying for many years the effects of various metals on bone, discovered an alloy they called Vitallium [(cobalt (65%) - chromium (30%) - molibdenum (5%)] and one of the first metals found to be completely electrically inert within body fluid, showing no signs of corrosion or pathologic changes to bone).  The material was first used to make plates and screws for fracture fixation, as was later adapted for femoral head prostheses. 

-In 1937: Smith-Petersen (first heard of this vitallium from his dentist who was using it for fillings, not from his orthopaedic colleagues) created the vitallium cup arthroplasty, which molded around the femoral head articular surface, and similarly struggled early fragmentation.

-In 1939: Dr. Venable worked with Dr. Harold Bohlman to design a Vitallium implant that replaced the femoral head (rather than cover it), similar to today’s hemiarthroplasty (it attached the vitallium femoral head to a small intramedullary nail). Dr. Venable was the first to perform a hemiarthroplasty for a hip fracture.

Around the same time, the Judet Bros (Paris, France) created a heat-cured acrylic femoral head for a similar hemiarthroplasty design that experienced high rates of early failure due to rapid wear and severe tissue reaction (considered the first cases of osteolysis). 

-In 1940s: Dr. Austin Moore (who was a surgeon at a psychiatric hospital in Maryland) worked further with Dr. Bohlman to improve the hemiarthroplasty design, and introduced the process of fixing the femoral head to a stem that would fit the proximal femur medullary canal

-In 1952: Dr. Moore and Bohlman introduced the stem that would allow for bone ingrowth (work was done at Austenal Labs which became Howmedica Inc. which became Stryker Corp.), and these produces were widely distributed, became legendary and are still used for some elderly femoral neck fractures. 

-In 1950: Dr. Fredrick Thompson created a Vitallium prosthesis, similar to the Moore type prosthesis, which used cement to fix the stem to the femur.  The prostheses developed by Dr. Moore and Thompson lead to acetabular erosion, highlighting the need to address both sides of the joint

-In 1950s: Ring, McKee-Farrar created the first Total Hip Arthroplasty, which was a metal-on-metal devise, which had too high friction and caused early loosening. 

-In 1950s: The modern version of a Total hip Arthropasty began with the works of Sir John Charnley, who transformed the procedure into a reliable and reproducible treatment for hip arthritis.  He is probably best known for pioneering poly liners and cement fixation.

-First to use cold-curing acrylic cement (polymethylmethacrylate, PMMA).

-“Low frictional torque arthroplasty”. Recognized that high forces on the implant was the major contributor to loosening.  In searching for low friction material combinations, he found that a stainless steel ball on polytetrafluoroethylene had a coefficient of friction similar to the normal joint.  He additionally decreased the femoral head size from 40 mm (which was the Moore type) to 22 mm (reducing the contact area and thus the frictional force).  Although decreasing surface area increases the pressure on the implant, he believed the key to longevity lay in reducing friction, and thus torque on the implant. 


metal shell of acetabular implant in total hip replacement


The current acetabular implant is a cementless 2-component cup.  A metal shell for bony ingrowth + a polyethylene liner (sometimes metal or ceramic) that snaps into the shell and  articulates with the femoral head. 

The modern acetabular implant is a "press-fit" design (it is not cemented).  The actual implant is about 1-4 mm smaller than the reamer of the same size (i.e. a size 50 implant is slightly smaller than a size 50 reamer).  This creates hoop stresses that hold the shell rigidly in place, allowing bone to grow into the shell.  Lets look into this process more.

The acetabulum is prepared via sequential reaming - removing the remaining cartilage and subchondral bone to create a bed of vascular cancellous bone that encourages bone ingrowth.  

If the target is cup size is 50, a surgeon will start smaller, ie 46 reamer, and work up to a size 49 or size 50 reamer.  If a size 50 reamer is the final one used for a 50 cup - this is called "line-to-line" fitting.  If a size 49 reamer is the final one used for a 50 cup - this is called “oversizing” the implant to increase hoop-stress on the implant.  Why oversize?  The hemisphere created by sequential reaming may not be a perfect circle if the surgeon moves the center-point of each ream, creating more of an oval, which decreases hoop stress.  Oversizing can compensate for imperfect reaming.  Alternatively, poor quality bone, as seen with avascular necrosis or osteoporosis, may reduce the viscoelasticity of bone, which in turn decreases its ability to generate hoop stresses, and therefore, oversizing may generate better hoop stresses (however, note that too much hoop stress will create a fracture, and poor bone is most susceptible).  

Hoop-stress create the press fit.  If the acetabular implant is undersized (ie a size 52 reamer for a size 50 cup), then stress is overly concentrated centrally (the implant doesn't fully touch the outer walls because reaming diameter is bigger), and there may be gaps between the component and bone at the periphery, preventing stability and proper ingrowth.  In contrast, if the component is too large for the prepared cavity (ie size 47 reamer for a size 50 cup), excessive stress is transferred peripherally with risk of acetabular rim fracture.  The center of the component also may not seat completely, also risking instability via ingrowth failure. 

note: hemispherical cups are “oversized” as described above (ie 50 cup prepared with a 49 reamer).  However, elliptical design cups are usually prepared “line-to-line” (ie 50 cup prepared with a 50 reamer), or even over-reamed by 1 mm.  This is due to the difference in geometry: you are using a hemispherical reamer but implanting an elliptical shell, this mismatch creates the requisite hoop stresses.

The metal component is highly polished on the inside and the outside is completely coated to allow for bony ingrowth (only about 30% ingrowth occurs).   The biologic fixation occurs by either bony ingrowth or bony ongrowth. The optimal pore size is between 100 and 400 μm.  Bone ongrowth occurs with a “roughened” (but not porous) surface.  This roughened surface occurs by grit blasting (a pressurized spray of aluminum oxide particles to produce an irregular surface at 3-8 μm depth, and a thickness of 50-150 μm), plasma spraying (apply molten metal in a argon gas environment), or hydroxyapatite coating (which is an osteoconductive calcium phosphate coating applied by plasma spray).  

note: newly developed highly porus metals (ie Trabecular metal) increases the friction against cancellous bone and improves initial stability (it also encourages rapid and extensive bone ingrowth). 

Many surgeons supplement press-fit fixation with screws to improve stability.  Retrieval studies show that most bone ingrowth occurs around these screws, highlighting their efficacy. 

Most metal shells are sized 40 to 70 mm (referring to their outer diameter), increasing in size by 2 mm.  Metal shells are ~ 5 mm thick to prevent fatigue fracture.  A shell can accommodate varying thickness of poly liner.  For example, a size 50 shell can be fit to match both a size 28 and size 32 femoral head.  The difference is the thickness of the liner. 

The most common size in women is 48, the most common size in men is 52.  

Historically cups were monoblock polyethylene (no shell) and were cemented into the acetabulum.  The poly was designed with backside grooves to improve the cement mantle.  However, studies showed cementing a poly cup leads to higher rates of aseptic loosening, particularly in long-term follow-up, because cement does great with compressive forces as seen in the femoral canal, but poorly with shear forces (which is what you see in the acetabulum).  Some studies suggest equivalent survival as compared to cementless cups with thick-walled poly (>5 mm) and preservation of subchondral bone.  A titanium shell that allows bone ingrowth was the solution for aspetic loosening of cemented cups, and while it appears to improve implant longevity, a poly liner with metal backed shell has unique problems.  

Disadvantages to metal backed cup. 

1. Need for a “locking mechanism” (fix the poly to the metal shell).  Its just one more thing that can fail, and there are reports of dislocated poly liners due to improper insertion or failure of the locking mechanism. 

2. “Backside wear”: Micromotion can occur between any two segments, no matter how well fixed they appear.  “Backside wear” refers to micromotion between the back of the poly liner and the metal-cut.  Many of the current designs have a polished inner poly shell to minimize the wear particle production, which can lead to osteolysis.

3. Increased bone loss. The metal-backed cup must be >5 mm thick to prevent fatigue fracture.  The poly must also be >5 mm thick to prevent fatigue fracture and early wear.  Thats a total of at least 10 mm thickness.  A single-component (ie metal only, or poly only) requires only 5 mm overall and thus takes up 50% less volume.  Adding the metal backing to a poly liner thus requires extra bone resection, unless you ream the same amount of bone and then make the poly smaller and use a smaller femoral head.  Yet many surgeons associate smaller heads with larger risk of dislocation.  Therefore, to preserve the larger head size, they will ream more native bone from the acetabulum.  Its follows the saying “robbing paul to pay peter”  (femoral head size vs. acetabular bone stock). One advantage to metal-on-metal design is that it can accommodate a larger femoral head because it’s a single component.

4) There is less concern about stress shielding in the acetabular component as compared to the femoral component, however it does occur with about 20% decreased density anteriorly (yet the clinical significance is less apparent).  This is particularly seen in monoblock cups made of cobalt chromium for metal-on-metal hips (as opposed to standard titanium shells which have modulus of elasticity that better matches native bone).


types of liner for acetabular implant in THA

Polyethylene liner (ie "poly") is the most common material type for the acetabular liner, however, ceramic or metal are also available. The Acetabular Cup provides the base for bony ingrowth, while the liner clips into the shell and articulates directly with the Femoral Head.  

The Liner Size refers to its inner diameter = the Femoral Head Size.  The poly must be at the very minimum 6 mm in thickness to prevent fracture. 

Femoral Head Size + Liner Thickness = Acetabular Cup size.  Liners come in varying thickness. For example a Size 32 Femoral Head can fit into an Acetabular Cup Size 48, 50, 52 etc etc because you can get a 16 mm, 18 mm, 20 mm etc thickness poly.  Importantly, there is variability in sizing between manufacturers.


Poly is a long-chain polymer that’s a tough biocompatible material, and the first iteration, polytetrafluoroethylene (aka Teflon), was introduced to THA by Dr. Charnley.  The original poly demonstrated high wear rates (about 0.2 mm/year) that lead to loosening from osteolysis.  There was minimal poly cross-linking, and it was sterilized in air (leading to high free radical production).

Interestingly, this loosening was initially attributed to cement failure, termed “cement disease”, and this misconception drove advances in press-fit technology.   Once the etiology of aseptic loosening was correctly attributed to conventional poly, poly was then seen as the leading barrier to long-term THA survival and thus multiple alternatives were developed, particularly the now notorious metal-on-metal implants. 

Yet advances in poly have continued over the decades.  The poly material improved to high-density polyethylene (HDPE) in the 1970s (wear 0.10 mm/year), and then to ultrahigh molecular weight polyethylene (UHMWPE) in the early 2000s, which has progressively demonstrated significantly less wear (<0.02 mm/year).

UHMWPE. There is a process to manufacturing UHMWPE.  The poly is subjected to higher radiation (5-10 Mrads) that breaks poly bonds and creates free radicals that bond with other free radicals on neighboring chains to form cross-links.  This is how cross-linking occurs.

The process initially creates instability, which can be problematic if the environment isn't closely regulated, for example, if there is oxygen lying around, then free radicals don't bond with each other to form cross-links but rather combine with oxygen, which propagates further free radical formation, and ultimately breaks down the poly to create "Oxidized PE".  Therefore, radiation is performed in the setting of inert gas to prevent oxidation of the poly.  Irradiating the PE in inert gas accomplishes two things: 1) it reduces wear by forming cross-links; and 2) it sterilizes the PE. There are techniques to sterilize the PE without cross-linking using ethylene oxide gas or gas plasma spray.

After the poly undergoes radiation in an inert gas, it is heated (via remelting, or annealing) to quench the remaining free radials (the heating allows free radicals form stable carbon-carbon covalent bonds).  This process has the side effect of decreasing the crystallinity, thus decreasing toughness and tensile strength of the poly.

Some second generation UHMWPE liners are impregnated with antioxidants (such as Vit. E) to further decrease free radical breakdown after implantation. 

Up to this point the poly is prepared as a solid tube of plastic.  It then needs to be shaped into the poly insert that’s implanted during surgery. There are a few techniques, although direct compression molding (implant made from a mold, no machining involved) creates the lowest wear.

In summary, its believed that the summation of advancements in poly manufacturing have decreased wear by 95% compared to conventional PE.  

Liner Design Types.

The standard poly liner is a neutral face hemispherical design to allow maximal range of motion (the goal is to provide a large implant ROM so that the THA falls within the motion circle allowed by native hip anatomy.

The poly can have a 10, 15, or 20° lip liner depending on the manufacturer.  The “lip” is placed in the region with the greatest risk for dislocation, to provide an additional few millimeters of clearance needed for the jump distance.  The poly can also be lateralized by 4 mm.  In this scenario, the liner has more material on the medial side, as opposed to the apex, thereby “lateralizing” the center of rotation of the hip.  This is the same as increasing offset, only it occurs on the acetabular side as opposed to the "high offset" stem.  This usually also results in adding length but it is negligible.  This liner can be used in cases of protrusio or revision cases when the goals are to increase stability by restoring tension on the soft tissues (particularly the abductor complex).  


The Femoral Head articulates with the Acetabular Liner.  The Femoral Head also forms a junction with the Femoral Stem at the Trunnion.  Technically there is motion at both ends of the femoral head (even though motion at the trunnion would ideally not occur) and thus both are important to consider as sources of wear debris.  The material of the femoral head can vary from the historical standard cobalt-chrome to ceramic to zirconium (which is a metal-ceramic hybrid).  The variation in femoral head material causes different levels of friction with the acetabular liner and generates different amounts of wear.  We discuss each of the femoral head materials below. The femoral head junction with the femoral stem trunnion has received a lot of attention lately as a current source of corrosion (see complications section).

Femoral head materials.

There are a number of materials for the femoral head, but the overwhelming majority of modern day THA utilizes a Ceramic or Co-Cr femoral head on a poly liner as the bearing surface.  The wear rates of these couples is extremely low with a slight advantage to ceramic-on-poly. 

Metal.   A Cobalt-Chromium (Co-Cr) femoral head is a good bearing material because it possesses low wear properties, while titanium alloy scratches easily leading to rapid polyethylene abrasive wear (titanium is a great material for the acetabular shell or femoral stem due to its modulus of elasticity).  The Co-Cr femoral head can articulate with a metal liner (hard on hard bearing) or with a polyethylene liner (hard on soft bearing). A Co-Cr femoral head initially has some rough spots (called "asperities") that cause accelerated wear ("run-in wear") in the first year (1 million cycles), however, these areas smooth out, and wear rates drop in subsequent years: initial wear is 0.18 mm/yr, then 0.10 mm/yr after wearing in.

When a metal head articulates with a metal liner (hard-on-hard), it forms a fluid-film layer during walking, which is a layer of lubrication between ball and socket, and it significantly decreases friction.  The fluid-film layer only exists with hard bearing surfaces because the material is so smooth (Co-Cr head surface roughness is only 0.01 μm, compared to poly roughness of 7.0 μm, while ceramic roughness is even better: 0.006 μm).  A larger head and high congruence between the head and liner maximize contact area and promote the fluid-film lubrication.

Complications.  Metal-on-metal wear can be a major complication if significant wear triggers an immune response (more in Complications Chapter). Wear particles are tiny (0.015 - 0.5 μm, yet the overall number of particles is significantly greater than poly wear, and it creates an adverse local soft tissue reaction (ALTR). 

Ceramic. An alumina-ceramic femoral head is a good bearing material because it is the smoothest material and therefore possesses the lowest wear properties (low abrasive, linear and volumetric wear).  The wear rate of ceramic-on-ceramic is virtually zero (<0.07 μm/year), and even the few particles generated are completely inert; the body mounts no immune response.  

Ceramic is hydrophilic, and absorbs moisture to promote fluid-film lubrication in ceramic-on-ceramic bearing.

Complications. While ceramic sounds like the wonder material, the long-term outcomes are nothing spectacular (86% survivorship at 18 years) [24].  Ceramic liners are reserved for the “ultra-young” patient (ie < 50 years old) because of unique complications: liner fracture and squeaking. 

Older designs had large pore size in the ceramic, this decreased strength, and led to ceramic fracture of 13% (a devastating complication, as the microscopic shards are near impossible to remove and contribute to significant third-body wear of the poly after revision - furthermore, despite best efforts, small shards remain and thus revision surgery requires the use of more ceramic because the density of ceramic makes it most resistant to third-body wear.  5-year survivorship is only 60% after revision for fractured ceramic. Using a ceramic head in the revision setting, placing it onto a retained stem,  requires a titanium jacket to prevent the roughened neck from causing a fracture). Yet, similar to the multiple iterations to improve poly, the modern ceramics undergo hot isostatic pressing to decrease grain size by 3x. The rate of fracture is now only 4 in 100,000 at this time (0.012%), with 40% fractures occur by 1 year, 75% by 3 years. Head size decreases fracture risk. Current ceramic heads do not have cobalt-chrome liners and thus decrease trunnionosis [25].  

note: oxidized zirconium is a new development (third generation ceramic), and is essentially a metal-ceramic hybrid (or “ceramizied metal”).  A metal alloy undergoes an oxidation process to create a zirconia ceramic surface (the outer layer of metal becomes ceramic).  This allows the surface to become harder with fewer bumps (ie low wear characteristics of a ceramic)  without the risk of fracture and chipping that remains a concern. This material is only FDA-approved for articulation with a poly liner.

Another deterrent to ceramics is “squeaking” (incidence is low: only 14 in 2138 THA, about 0.6%). Its an audible squeak heard with every step, typically develops more than a year after implantation, and annoys patients enough that many undergo revision surgery. Its cause remains unknown, it may be associated with stripe wear due to implant malpositioning (a vertical cup > 55 degrees, causing edge loading which disrupts the lubrication of the bearing.


Size is another important technical consideration.  20 years ago, the average femoral head size was 22 mm.  Today the average size is 32 mm.  That’s almost 30% bigger.  Does this mean that bigger is better? Most joint surgeons agree that there are both benefits and risks to a larger head size. 

Benefit. A larger head increases stability for two reasons.  A larger head increases the head-neck ratio (diameter of the femoral neck vs the femoral head… the neck diameter never changes).  A larger head-neck ratio means a larger arc of motion before impingement (eventually the neck will impinge on the rim of the socket). Because impingement causes the head to lever out of the socket and dislocate, a larger head decreases the chances of impingement, thus increasing stability.  However studies have found that 32 mm heads and larger effectively prevented impingement between components (without significant added benefit from heads larger than 32 mm). Furthermore, if the femoral neck was made trapezoidal instead of circular, impingement decreased significantly (while a skirted head dramatically increased impingement). 

Additionally, a larger head stabilizes the joint because it increases the jump distance (Jump Distance = the radius of the femoral head, and represents the distance that the head must travel (sublux) before full dislocation occurs).  If the head is larger, then there is a greater distance before dislocation. 

Disadvantage. Increased debris formation for two reasons.  A larger head increases volumetric wear of the acetabular liner and thus increases the amount of poly debris.  A larger head also increases the head : neck ratio, which is great for arc of motion but it also places more stress across the head : neck junction (aka the trunnion) and is associated with corrosion (“trunnionosis”) which produces a similar reaction to the highly-undesirable Metal-on-Metal wear. 

Additionally, a larger head requires a larger poly liner, which can be achieved only by making a standard poly liner thinner (a thinner poly has a higher risk of breaking or wears out to a critical level faster); or by using a larger acetabular shell (which means taking away extra bone stock).

stem implants


important landmarks of the femoral stem

PRESS-FIT STEM (cementless)

The cementless stem obtains initial stability via “press-fit” allowing "bone ingrowth" of the porous-coating by six weeks post-op.  

1. rigidity is the short-term goal.  

The key to fixation is immediate implant rigidity .  A good “Press-Fit” is direct contact between the implant and weigth bearing cortical bone (gap <50 μm) and no micromotion (<30 - 150 μm, over which leads to fibrous ingrowth which is painful and unstable).   

The stem is designed so that it is slightly larger than the femoral canal that was prepared through sequential broaching, and by wedging in a slightly "oversized" stem, the hoop stress prevent motion.  

The alternative is "line-to-line" preparation of the femoral canal, the canal and the stem are a perfect match in size, however, the rough coating of the stem creates enough friction to prevent motion (aka "scratch fit" or "interference fit").

The variability of femoral canals (ie Dorr Classification) prevents a single stem design from truly standing out.  The design type determines the area of the femoral canal where fixation occurs and the surgical technique required for implantation.   

2. biologic fixation is the mid-term goal.  

The goal is for bone to incorporate into the implant.  Therefore, the implants are typically titanium alloy so the implant modulus of elasticity is the closest match to bone. There is either bony ingrowth or bony ongrowth. Bone growth is circumferential to prevent communication between the joint space and the distal aspect of the component (which is called a large "effective joint space" and can occur when there is incomplete proximal coating). 

Bone ingrowth occurs with porous coating, achieved by heating the metal (which can decrease the fatigue strength).  The optimal pore size is between 50 - 150 μm), the deeper the pores, the better the ingrowth, and overall porosity 40 - 50%.  

Bone ongrowth occurs with a “roughened” surface (divots not pores).  This roughened surface occurs by grit blasting (a pressurized spray of aluminum oxide particles to produce an irregular surface at 3-8 μm depth, this depth = the "roughness" and greater roughness = greater fixation). Plasma spray (applying molten metal in a argon gas environment) also creates roughened surface. The depth of pores and depth of divots are both directly proportional to strength of fixation.  Because pores are deeper than divots, its clear that ingrowth gives better fixation than ongrowth per mm^2.  To compensate, the surface area of bone ongrowth must be greater (and its used almost only for stems, not cups).  Furthermore, the Bone ongrowth technique requires press-fit technique (cannot use scratch fit).

difference between bone in-growth and bone on-growth for press fit femoral stem

note: bone ongrowth can also occur with stems containing splines (longitudinal grooves that travel along the diaphyseal portion of the stem)

Hydroxyapatite coating is an osteoconductive calcium phosphate coating applied by plasma spray.  It is shortens the time to biologic fixation.

3. BALANCED STRESS distribution across the femur is the long-term goal of cementless stems.  

The key to longevity is equal distribution of stress across the prosthesis.  

Uneven stress causes “stress shielding” the areas that do not experience stress.  Stress shielding is the resorption of bone over time because the bone is not functioning in its normal capacity, and thus answers in the affirmative to the question: “is it true that if you don’t use it, you lose it?”  In contrast, areas that experience excess stress are painful and prevent bone ingrowth. Titanium alloy is preferred because its lower modulus of elasticity is most compatible with cortical bone (shares the stress) thus decreasing thigh pain (yet one study [26] compared identical design stem in titanium and cobalt-chrome and showed all thigh pain and stress shield was only related to larger stem size because the radius is such a powerful determinant of rigidity, remember its r4).  

GEOMETRY OF STEMS. Multiple designs attempt to achieve the aforementioned 3 Goals of uncemented hip arthroplasty. 

femoral stem designs

Flat Tapered, "Single-Wedge", "Blade". This design refers to an approach for achieving early fixation by wedging a rectangle into a circular hole.  The intentional mismatch of geometry achieves good torsional stability.  Fixation occurs by medial-lateral engagement and by the 3-point fixation along the length of the stem.  Its broach only. There is no reaming because there isnt significant distal engagement as the stem tapers in the medial-lateral plane.  The minimal distal contact between implant and diaphysis helps to prevent proximal bone resorption (known as "disuse osteopenia" or "stress shielding" which is an issue with the cylindrical stems that engage the diaphysis).  The stem often has a recessed shoulder to reduce the need to resect bone laterally.  

Every implant company makes a stem with this design.  Smith & Nephew: Anthology, DePuy: Tri-lock , Stryker: Accolade II , Biomet: Taperlock, Zimmer: M/L Taper.  

Dual Taper, "Elliptical".  This design refers to an approach for achieving early fixation by completely filling the metaphyseal canal and circumferentially engages the femoral cortex proximally.  The technique is often referred to a "Fit-n'-Fill" technique requiring both reaming and broaching.  The implant tapers in all planes and is overall more substantial in volume as compared to a single-wedge design.   

These designs are usually made for cementless and cemented option.   Every implant company makes a stem with this design.  Smith & Nephew – Synergy; DePuy – Summit;Stryker - Secur-Fit, Biomet - Integral, Zimmer - Versys.  

Cylindrical.  This design refers to an approach for achieving early fixation by engaging not only the metaphysis but also the diaphysis for ingrowth.  The cylindrical model engages the diaphysis because there is no taper (unlike the aforementioned examples).  It was developed to address patients with poor metaphyseal bone quality that put them at risk for failure to achieve initial stability or later ingrowth.  The disadvantage of the stems is that ingrowth is rarely uniform across the entire stem, and the majority occurs distally, which then assumes the majority of the stress leading to "stress shielding" in the proximal metaphysis, which leads to significant bone loss.  This design is used for primaries, although it has been replaced by the two aforementioned designs in most centers.  This design is rather preferred for revision surgery.  

Most implant companies make a stem with this design.  Smith & Nephew - Echelon, DePuy – AML, Solution , Stryker - Secur-Fit Max, Biomet – Mallory-Head, Zimmer - Epoch

revision femoral stem designs

Wagner-Type. This design refers to an approach for achieving early fixation by engaging the diaphysis for ingrowth.  The Wagner-type design is a tapered cylinder (axial stability) with splines to grip the cortex circumferentially (rotational stability).  The initial design was monoblock (stem is one unit), however, many of the popular current designs are "Modular", meaning there are two parts to the stem that allows the surgery to adjust stem height once the diaphyseal portion has been implanted.  A Modular Stem is beneficial for a few reasons.  For one, you can adjust the degree of anteversion to help reduce dislocation risk.  Additionally, because these stems are often used in the face of significant bone loss, there is some subsidence of the stem when its inserted (meaning: the line you use to measure depth when you are reaming the femoral canal does not always equal the height of the actual stem once its inserted).  Some of this depends on the steepness of the taper.  A 1 degree taper is a shallow transition and may subside more than a 3 degree taper, which is sharper and thus prevents some subsidence.  The standard taper is a 2 degree, which some variation between companies.   

Most implant companies makes a stem with this design.  Smith & Nephew - Redapt, DePuy - Reclaim, Stryker - Restoration, Biomet - Arcos , Zimmer - ZMR or Wagner SL.  

Bone Preserving. This design refers to an approach for metaphysial for ingrowth that attempts to preserve bone by performing a high femoral neck cut.   The implant is smaller than a standard flat taper stem, and looks like a small "chili pepper", but obtains fixation similar to the single wedge taper design that relies on 3-point fixation. There is greater variability in the femoral neck person to person than in other areas of the proximal femur.  Therefore, the problem with this design is that multiple iterations of one stem (varus, valgus, etc) are needed to account for these variations.  

Some implant companies makes a stem with this design.  Zimmer - Fitmore, Biomet - Microplasty.


The cemented stem was the original THA system (Charnley stem was stainless steel). 

The implant is typically highly polished Cobalt-Chrome. The high-polish seems counterintuitive (“don’t you want the cement to interdigitate into the implant?”), but remember that cement is strongest in compression, weakest in resisting shear forces.  There is no interdigitation because this would emphasize shear stress, while the polished design allows the stem to compress into the cement mantle (it subsides slightly, less than a milimeter, in the first 24 hours, giving added strength).  A second reason that cemented stems are smooth is that if debonding does occur, a rough surface will produce more wear particles than a smooth surface, thus accelerating osteolysis and loosening (as seen in the Exeter Stem which had a matte surface). 

Collars are commonly associated with cemented stem designs as a feature to determine depth of insertion, prevent subsidence, and add stress to the calcar region to prevent stress shielding.  The collar however does not prevent subsidence in the long run in a meaningful way as compared to collarless stems.

collared stem vs. collarless stem:  As discussed above, the calcar is at greatest risk from stress shielding. Could a collared stem increase stress to this region and counteract some of the disuse osteopenia. Studies demonstrate that proximal medial bone strain is 65% normal in a collarless press-fit stem, and 70-90% normal in a collared press-fit stem (however, loose fitting stem cause strains greater than a normal femur).  A cemented collar also has the theoretic benefit of increasing stress transfer to the calcar region. A cemented collar also helps control height during insertion (prevents subsidence while the cement hardens). Others argue that a collar effectively prevents subsidence of the component during the initial stage after implantation which relies purely on the press-fit before ingrowth has occurred.  Others argue however, that such subsidence is desirable, as a wedge fit prevents the micromotion that causes fibrous-ingrowth and persistent motion.  If the collar engages the neck cut before the metaphyseal portion of the component engages, there is increased risk for fibrous-ingrowth and persistent micromotion. 

The stem itself should occupy 80% of the medullary canal.  Central stem placement minimizes the chance of uneven cement mantle, preventing areas of thin cement. The target Cement Mantel is 4 mm proximally, 2 mm around the stem distally (providing sufficient thickness to prevent cement fracture). A cement centralizer affixed to the distal stem before implantation will improve the symmetry of cement mantle.  The ideal length of the stem is around 13 cm (which was used by Charnley) because it allows for sufficient distal cement mantle, while a longer stem is problematic due to the native anterior femoral bow and canal narrowing near the isthmus, which together prevents a reliable cement mantle. The high modulus of elasticity in Co-Cr (in contrast to the titanium used for porous coated cementless stems) decreases proximal cement mantle stress. 

Failure. When a cemented component fails, it typically fails at the component-cement interface because of debonding and then cement mantel fracture.  One attempted solution was to pre-coat the stem implant with cement (the Harris Stem), however, as we now know, this design lead to a high failure rate because it emphasized shear stresses and only made matters worse at the component-cement interface.

CEMENT Overview

PMMA (polymethylmethacrylate) is the bone cement used it TJA. PMMA was discovered in 1843, the first medical application of PMMA was in dentistry, and in 1945 the Judet Bros were the first to use it in orthopedics, as a femoral head prosthesis.  However, its modern use (as a cement, not an implant material) was first used by Haboush in 1953, and then famously by Charnley in 1970 where it revolutionized total joint arthroplasty. 

PMMA it goes by different names, such as Simplex P (Stryker brand – white color), Palacos R (Heraeus brand – green color because of chloraphyll additive) and CMWI (DePuy brand).  These brand names use variations of the same base ingredients. There is an international governing body that actually regulates the standardization of biomaterials, and companies are required to comply to the mechanical and working properties of the compound (ie ultimate tensile strength, bending strength, compression strength, shear strength and fracture toughness).  So whats the difference between cement brands?  The final product is the same.  Yet as the cement hardens, it goes thru three phases.  The sticky phase, the working phase, and the hardening phase.  Different cement brands have different viscosities. A high viscosity cement has a very short sticky phase and a long working phase.  A low viscosity cement has a long sticky phase and short working phase.

Basic Properties. PMMA is a polymer. Polymers have weak non-covalent bonds between side chains that break under constant strain, which leads to gradual deformation (creep) and stress reduction.  This material behavior is similar to both an elastic solid and a viscous liquid thus giving the commonly used name “visco-elastic”.  This property is important clinically because it indicates that implant subsidence can occur within the cement mantle, particularly during the initial few weeks after THA as the PMMA actually continues to polymerize for a few weeks in vivo (and creep is highest in fresh cement as there are more side chains that can break, leading to deformation).  Furthermore, antibiotics increase the creep of cement, and actually water (which bathes the cement in vivo) increases the creep of cement.  Interestingly, while the varying cement brands are uniform in their static property, there is variation in their creep due to small chemical differences.  Creep is a complex process that still requires more study.

The static properties of PMMA.  Poor tensile strength (25 MPa), moderate shear strength (40 MPa), and is strongest in compression (90 MPa). The modulus of elasticity for PMMA is 2400 MPa (10x less than cortical bone and 100x less than the metal implant), therefore it’s relatively soft material sandwiched between two rigid materials.

Cement Technique. PMMA is a polymer but its not the only chemical contained within the cement as its prepared. As you may know from prep on the back table, cement is made by combining a liquid and powder. 

The powder contains the polymer (polymethylmethacrylate), as well as barium sulphate (or zirconium dioxide), which act as radiopacifiers, a dye (ie chlorophyll) to distinguish the cement from bone, and it also may contain antibiotics.

The liquid component contains the monomer liquid (methylmethacrylate) which allows for polymerization at room temperature.  The chemical DMPT (n,n-dimethyl-p-toluidine) initiates cold curing at room temp., benzyol peroxide reacts with DMPT to catalyze polymerization, and hydroquinone (stabilizer to prevent premature polymerization).

Once the liquid and powder components are combined, they are mixed under vaccum suction to reduce the porosity and thus increase the tensile fatigue strength.

Bone preparation.  Cleaning the bone (via lavage) and drying the bone (using packing or suction) maximizes cement interdigitation within the bone.  Greater surface area of cement fixation improves fixation.  Similarly, pressurizing the cement (using the cement gun and cement restrictor) enhances interdigitation of cement.

Stem centralization.  A cement mantel is only as good as its thinnest portion (like a chain as strong as its weakest link).  A stem centralizer can help obtain a uniform cement mantle to ensure the minimum 2 mm cement surrounding the implant to minimize the risk of cement fracture. 

Stem Stiffness.  In contrast to a press-fit stem where titanium is the preferred metal because it matches the modulus of elasticity, the cemented stem performs best when the metal is rigid and thus decreases stresses on the cement mantle.  Cobalt-chrome or stainless steel are preferred materials.

Antibiotic Cement [4]. The practice of adding antibiotics to cement to either prophylax against infection or treat a periprosthetic infection began in 1970 by Buchholz and Engelbrecht [5]. An antibiotic is either added at the time of surgery or a commercial-antibiotic-impregnated-cement can be purchased off the shelf.  In the USA, antibiotic cement is FDA approved for the use of revision surgery for periprosthetic infection but is not approved for use in primary TJA. The primary concern with using antibiotic cement for primary TJA is the development of bacterial resistance.  Hope et al [6] demonstrated that 90% of s.aureus infections were resistant to gentimicin in the antibiotic-cement group, while only 16% were resistant if normal cement was used.  However, studies demonstrate a statistically significant reduction in deep infections for primary TJA using antibiotic bone cement[7], and thus the AAOS advices for its use as prophylaxis only when considerable risk factors exist [8].  However, a randomized control trial in Sweeden, comparing systemic antibiotics (ie perioperative ancef) compared to antibiotic impregnated cement, demonstrated no significant difference for infection rate.  A survey of surgeons in 2003 indicated that 12% always use antibiotic cement (primaries and revisions), 69% never used it, and 19% sometimes use it.

The properties of antibiotic release from cement varies based on the type and concentration of antibiotic. The effect of antibiotics depends on concentration, as higher concentration within the cement enables higher concentrations within the joint over a longer time [9].  Local levels of antibiotic are generally significantly higher than the minimum inhibitory concentration (significantly higher than levels obtained with intravenous antibiotics), but the ideal concentration to treat an infection is less clear…is higher concentration always associated with better clinical treatment?

Gentamicin is historically [10] the most commonly used antibiotic for cement because its family (aminoglycosides) are very thermally stable (remember that cement heats to >100° C during exothermic reaction), while also providing broad coverage and is water soluble (must be water soluble to diffuse to surrounding tissue).  2 grams of genticmin added to 40 g of the polymer has no significant impact on strength, while >4.5 g appears to significantly affect strength. This 5% concentration of antibiotics is an acceptable compromise between antibiotic strength and cement strength, although surgeons have reported increased concentrations without seeing clinically significant complications. Surgeons have reported from the Mayo Clinic, using 4 g vancomycin per 40 g batch of cement and 4.8 g gentamicin per 40 g batch of cement for their antibiotic spacers without observing toxicity in patients or cement failure.

The compressive strength to failure does not appear as significantly impacted by antibiotics as fatigue failure.   Vancomycin concentrations, in contrast, have a more linear impact on the static properties of the cement.  When 1, 2, or 3 g of vanco is added to cement, the cycles to failure is 90%, 70%, 50% compared to normal cement, and 1.2 or 2.4 g of tobramycin cement failed at 80% and 60% cycles compared to normal. Rifampin changes the chemical reaction of cement and slows setting-time from minutes to days, and therefore cannot be used.  Furthermore, the compressive strength of cement can decrease by 40% if antibiotics are added as a liquid or if they are added to the monomer.

Multiple antibiotics in cement improves antimicrobial coverage, and may provide synergistic effects with regards to bacteriacidal effects and elution into the joint [11]. 

Other antibiotics have been used with varying success.  Penicillins are heat-stable, however, allergy to PCN is so common that its generally avoided.

Concentration of antibiotics. On average only 5-8% of the antibiotic contained within the cement is released.  Studies on tobramycin and vancomyicin suggest elution occurred over a 9 week study period, although peak levels occurred at 18 hours after implantation, and were 5x higher than levels after just 72 hours [12]. Yet, despite this seemingly rapid elution, vancomycin, due to its higher molecular weight, elutes less efficiently than aminoglycosides and therefore elutes for a longer time.

Palacos cement appears to be most efficient in releasing antibiotics, with CMW-1 releasing 24% less tobramycin and 36% less vancomycin.  Palacos similarly eluted antibiotics better than Simplex cement. 

7. Revision - Acetabulum

Bone loss is a major challenge with revision THA (particularly cases of osteolysis and infection). Bone loss means less bony coverage of the cup which means less surface area for ingrowth which means higher risk for failed fixation of the cup.  

Bone loss can also compromise initial stability.  A primary THA utilizes hoop stress of an intact acetabular rim to give stability for a press-fit implant (the press-fit is the primary source of implant stability until ingrowth occurs around 6 weeks).  The standard titanium cup designs require micromotion < 50 μm, and about 50% contact between the implant and viable bone for successful ingrowth [11] .   New porous metals, such as trabecular metal, have been developed for the revision setting where initial stability and bony ingrowth is a challenge.  This technology increases friction between the implant and bone, increasing initial stability despite less bony contact, and it encourages extensive rapid bone ingrowth [12].

Bone defects can be described as “segmental” or “cavitary”.  “Segmental” refers to defects of the acetabular rim, which compromises stability, because only an intact rim can withstand the hoop stresses that stabilize a press-fit cup.  “Cavitary” refers to volumetric bone loss where the rim remains intact.  Both however decrease coverage of the implant [13].

X-rays are used to determine the type and location of bone loss. This is achieved by looking for lysis at the ischium, teardrop and determining if the hip center has migrated superior, medial or lateral.   Judet views ("O.A.K" Obturator Oblique shows Anterior "K"olumn, Posterior Wall, Iliac Oblique shows Posterior Column, Anterior Wall) and a CT scan of the pelvis can be used to further characterize bone loss.  

- Superior migration of the hip center indicates bone loss at the dome.  If you identify superior migration, then determine if its Superior + Medial (indicating anterior column > posterior column bone loss) or Superior + Lateral migration (indicating posterior column > anterior column loss).

-Isolated medial migration ("protrusio") of the hip center is a defect in the anterior wall identified on x-ray as a break in the ilioischial line ("Kohler's line"). 

- Lysis.  Osteolysis of the ischium reports posterior column/posterior wall involvement.  Osteolysis of the teardrop indicates inferior bone loss.  Break in ilioischial line indicates anterior column bone loss. 

The Paprosky Classification was established to further characterize the severity of bone loss, which in turn guides treatment (different grades of bone loss require different implants to achieve stability) [14]. Lets look first at the Classification, and then at the corresponding treatment options.


Revision options are primarily based on Paprosky Classification. The amount of supportive acetabular rim present primarily determines if a hemispherical implant can be used in reconstruction. 

Type 1. Standard hemispherical cup is sufficient to get stability because the rim is completely supportive.

Type 2. Stable Rim but only partially supportive, meaning it may have some disruption  (segmental defect) without compromising overall stability. The key is that migration is not significant, and often a hemispherical cup is sufficient to achieve stable fixation, although an augment may be used to improve contact area for ingrowth.  

Type 3. The hallmark of Type 3 is an unsupportive rim. Thus a hemispherical cup cannot be used to achieve initial fixation.  This type is then separated into A and B.  

Type 3A is a posterior column/wall defect. The hip center migrates “up and out” (superior lateral migration > 3 cm).

Type 3B is an anterior column/wall defect (which has a higher association with pelvic discontinuity). The hip center migrates “up and in” (superior medial migration > 3 cm).  In a Type 3B, the surgeon must first determine if pelvic discontinuity is present before addressing the acetabular defect (otherwise the acetabulum will continue to open up and the cup will never achieve stable fixation).

If discontinuity is present, the options are a cage, a cup-cage, a custom triflange, or a distraction technique. 

Revision Cups

▪ Augments on a Hemispherical cup is required when the rim is not intact.  Titanium augments improve surface area of bony contact for ingrowth. Augments are becoming more popular as multiple configurations are available, such as the “flying buttress” for superior placement, or a “dome” augment for rim defects anywhere, and also the “footing” augment for medial defects.  [15]

▪ Triflange cup can also be used when the rim is not intact. Some cases of pelvic discontinuity can be addressed with a Custom Triflange cup, modeled based on CT imaging of the pelvis.  The implant contains locking screws to the ischium (inferiorly) and ilium (superiorly) for initial fixation, and its custom shape allows for sufficient bony contact for long-term stable fixation.

▪ Jumbo Cup.  In cases of large cavitary defects (contained lesions with < 50% bone contact), a Jumbo Cup is used (defined as outer diameter >66 mm). The Jumbo Cup is coated with highly-porous metal (ie Tantalum or Trabecular Metal) and provides better fixation via improved biocompatibility for host osteoblasts to ingrow.  These highly-porous metal additionally have a high coefficient of friction, and elasticity to give better initial fixation [16] [17].  

▪ Antiprotrusio Cages (“cage”) were used historically as the workhorse for pelvic discontinuity and to treat large contained lesions. The cage spans a large bone defect, and stability is achieved through proximal screw fixation to the ilium (superiorly) and engaging the ischium (inferiorly). A poly liner is then cemented into the cage. Because cages never achieve biologic fixation, there is a risk of fatigue failure [18], with failure rates around 20-30% at 5 years.  The cage fails via abduction pull out.

▪ Cup-Cage Concept is way to combine the benefits of both jumbo cups and cages [19].  The Cup-Cage technique inserts a "jumbo" trabecular metal shell (>60 mm) into the defect and adds fixation with multiple holes thru the cup into the acetabulum.  Then a cage (described above) is inserted on top of the cup and fixation is achieved with screws into the ilium and a slotted flange into the ischium.  The cage further stabilizes the acetabulum to minimize micromotion and promote ingrowth. A poly liner is then cemented into the cage.  The advantage of this technique is that it provides excellent immediate fixation (via cage) while also allowing for long-term biologic fixation (via jumbo cup). Good outcomes are reported for Cup-Cage technique [20]. It can also be used in Type 3A and 3B defects (with and without pelvic discontinuity) which still have significant bone loss making long-term fixation a challenge [21].  

Beyond bone loss, Instability is major challenge with revision THA, and it can be partially addressed through the choice of implants (for full discussion see Instability, in the Complications Section).  The femoral head can be upsized to increase jump distance and arc of motion before impingement.  The poly liner can provide increased version or increased jump distance, or it can be constrained.  

8. Revision - Femur

Bone loss is a major challenge with revision of the femoral implant.  In the face of significant bone loss, achieving fixation via press fit technique requires different implants, depending on bone loss severity.  

Geometry of Revision Stems. Revision stems are extensively coated (entire stem allows for ingrowth) because bone loss primary occurs in the metaphyseal region  after explanting a prior stem.  Many of these stems are modular to allow the surgeon to adjust length, offset and version.  To prepare for implantation, these implants require proximal broaching and distal reaming. 

Extensive-Coated Stem: Biologic fixation occurs in the metaphysis and diaphysis.  A well-fixed stem will have a "spot weld" at the tip of the stem.  The benefit of extensive coating is that metaphyseal bone loss is common in revision cases, and therefore is not reliable site for biologic fixation.  Diaphyseal fixation avoids this problem, however, because all of the force is transmitted to the area of fixation, the bone proximal to the fixation (ie the metaphyseal bone) will undergo stress shielding (it does not experience stress and therefore the bone will be resorbed, "use it or lose it"). The stem geometry also contributes to the degree of stress shielding.  Because a stiffer material will bear more stress, the stiffer stem will create more stress shielding.  The stiffest type of extensive-porous stem is Cobalt-Chrome (not titanium), round, solid (no flutes, no slots, no splines), and a large diameter (> 16 mm). 

Wagner Design: achieve more of a metaphyseal-diaphyseal junction fixation, tapered with a round conical geometry (some have cutting flutes for rotational stability).  The more distal fixation allows for use in revision cases.

Trabecular Metal is a highly porous metal which increases friction against cancellous bone to improves initial stability and encourages rapid and extensive bone ingrowth.  

The Paprosky Classification was established to grade the severity of bone loss, which in turn guides treatment (different grades of bone loss require different implants to achieve stability).  [1, 2].  

Revision based on Paprosky Classification

Type 1:  There is sufficient proximal bone to support any implant and therefore a primary stem (metaphyseal ingrowth) is sufficient (using a double wedge taper design which provides the best proximal fill), altought there may still be compromised rotational stability and thus many surgeons elected for a fully porous stem (diaphyseal ingrowth).  

Type 2: there is insufficient metaphyseal bone stock to support the stem, thus the implant must have some diaphyseal porous coating for ingrowth.

▪ Extensively Porous-coated Cylindrical stem (also referred to as the “monobloc” or diaphyseal porous-coated).  Proximal fixation isn’t possible, yet there is good diaphyseal bone for fixation, by extending the porous coating distally, ingrowth will occur in the diaphysis where there is bone.  This design is successful (3-5% revision rate at 15 years). [3] [4]

▪ Porous-coated proximal sleeve that is coupled with a smooth stem containing splines for rotational stability (the S-ROM type by DePuy or the Emperion by S&N).  This design requires enough metaphyseal bone stock for ingrowth, however, the modularity between proximal and distal components allows for management of significant mismatch in bone stock between the metaphysis and diaphysis [5].  The proximal “sleeve” contains a standard or a calcar-replacing option.  The sleeve has “step cuts” to convert shear force into compressive force.  This design promotes proximal ingrowth, it reduces the risk of stress shielding as seen in fully-porous coated stems. This design is best for revising undersized implants or early loosening.  The distal stem component is variable in length and can be curved to follow the natural femoral bow.  It contains splines for rotational control and a coronal split (“slot”) to reduce stiffness (and thus reduce thigh pain).  The Emperion by S&N has a “bullet tip” to further decrease pressure on the cortex and reduce thigh pain (it contains porous-coat plus Hydroxyapatite coating).  The biggest concern with the S-ROM type device is the extra modularity which enables micromotion at this additional joint (stem-sleeve) allowing for metal debris.  This micromotion can be mitigated in part by increasing the stem diameter to fully engaging the distal femur. 

Type 3A: no metaphyseal bone stock, some deficient proximal diaphyseal bone stock.  Requires diaphyseal fixation.  While standard fully coated stem is an option, good results have also been shown with the Modular stem designs.

▪ Extensively Porous-coated Cylindrical stem (also referred to as the “monobloc” or cylindrical diaphyseal porous-coated cobalt-chromic stem).  

▪ Porous-coated proximal sleeve that is coupled with a smooth stem containing splines for rotational stability (the S-ROM type by DePuy or the Emperion by S&N).  This cannot however be used for Type 3B [6].  

▪ Wagner Modular Stem.  This modularity separates the two goals of surgery, and allows the surgeon to work on them separately.   All fixation is achieved distally (in contrast to S-ROM), with the goal of bypassing the proximal deficient bone. Then the proper length, offset, and version is achieved proximally.  Then the two are connected.  Basic design is Tapered, Fluted Cylindrical Stem. Axial stability is achieved by wedging the taper.  Rotation stability is achieved with the flutes.  Its made of titanium to reduce stress shielding.  This design is the work-horse of revision THA in Europe for the past two decades before making it to the US [7]. Good clinical survivorship [8] [9]. 

Type 3B: even less diaphyseal bone stock (< 4 cm intact diaphysis) yet the femoral isthmus remains supportive. The Modular stem designs are the best option. 

▪ Wagner Modular Stem [10] is cylindrical and tapered.  

Type 4: without any supportive diaphysis, there is no way to obtain initial press-fit to allow for bony ingrowth.  Therefore these cases are incredibly challenging and typically require a total femoral replacement or APC (allograft-prosthetic composite).


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2.         Hamadouche, M., et al., Alumina-on-alumina total hip arthroplasty: a minimum 18.5-year follow-up study. J Bone Joint Surg Am, 2002. 84-A(1): p. 69-77.

3.         Kurtz, S.M., et al., Do ceramic femoral heads reduce taper fretting corrosion in hip arthroplasty? A retrieval study. Clin Orthop Relat Res, 2013. 471(10): p. 3270-82.

4.         Joseph, T.N., A.L. Chen, and P.E. Di Cesare, Use of antibiotic-impregnated cement in total joint arthroplasty. J Am Acad Orthop Surg, 2003. 11(1): p. 38-47.

5.         Buchholz, H.W. and H. Engelbrecht, [Depot effects of various antibiotics mixed with Palacos resins]. Chirurg, 1970. 41(11): p. 511-5.

6.         Hope, P.G., et al., Deep infection of cemented total hip arthroplasties caused by coagulase-negative staphylococci. J Bone Joint Surg Br, 1989. 71(5): p. 851-5.

7.         Parvizi, J., et al., Efficacy of antibiotic-impregnated cement in total hip replacement. Acta Orthop, 2008. 79(3): p. 335-41.

8.         Jiranek, W.A., A.D. Hanssen, and A.S. Greenwald, Antibiotic-loaded bone cement for infection prophylaxis in total joint replacement. J Bone Joint Surg Am, 2006. 88(11): p. 2487-500.

9.         Masri, B.A., C.P. Duncan, and C.P. Beauchamp, Long-term elution of antibiotics from bone-cement: an in vivo study using the prosthesis of antibiotic-loaded acrylic cement (PROSTALAC) system. J Arthroplasty, 1998. 13(3): p. 331-8.

10.       Wahlig, H., et al., Pharmacokinetic study of gentamicin-loaded cement in total hip replacements. Comparative effects of varying dosage. J Bone Joint Surg Br, 1984. 66(2): p. 175-9.

11.       Penner, M.J., B.A. Masri, and C.P. Duncan, Elution characteristics of vancomycin and tobramycin combined in acrylic bone-cement. J Arthroplasty, 1996. 11(8): p. 939-44.

12.       Penner, M.J., C.P. Duncan, and B.A. Masri, The in vitro elution characteristics of antibiotic-loaded CMW and Palacos-R bone cements. J Arthroplasty, 1999. 14(2): p. 209-14.

13.       Khanuja, H.S., et al., Cementless femoral fixation in total hip arthroplasty. J Bone Joint Surg Am, 2011. 93(5): p. 500-9.

14.       Bloebaum, R.D., et al., Postmortem analysis of bone growth into porous-coated acetabular components. J Bone Joint Surg Am, 1997. 79(7): p. 1013-22.