Nadim James Hallab
Department of Orthopedic Surgery, Rush University Medical Center, Chicago, IL
Biologic-reactivity to implant-debris is the primary determinant of long-term clinical performance. The following reviews: 1) the physical aspects of spinal-implant debris and 2) the local and systemic biologic responses to implant debris.
Methods included are: 1) gravimetric wear analysis; 2) SEM and LALLS; 3) metal-ion analysis; 4) ELISA, toxicity testing, patch testing; and 5) metal-lymphocyte transformation testing (metal-LTT).
Wear and corrosion of spine-implants produce particles and ions. Particles (0.01–1000 μm) are generally submicron (<1 μm). Wear rates of metal-on-polymer and metal-on-metal disc arthroplasties are approximately 2–20 and 1 mm3/yr, respectively. Metal-on-metal total disc replacement components have significant increases in circulating metal (less than 10-fold that of controls at 4 ppb-Co and 3 ppb-Cr or ng/mL). Debris reactivity is local and systemic. Local inflammation is caused primarily by ingestion of debris by local macrophages, which produce pro-inflammatory cytokines TNFα, IL-1β, IL-6, and PGE2. Systemic responses associated with implant-debris have been limited to hypersensitivity reactions. Elevated amounts of in the liver, spleen, etc of patients with failed TJA have not been associated with remote toxicological or carcinogenic pathology to date. Implant debris are differentially bioreactive. Greater numbers are pro-inflammatory; the smaller-sized debris are more bioreactive by virtue of their greater numbers (dose) for a given amount of implant mass loss (one 100-μm-diameter particle is equivalent in mass to 1 million 1-μm-diameter particles). Elongated particles are pro-inflammatory (ie, aspect ratio of greater than 3). Metal particles are more proinflammatory than polymers, ceteris paribus.
Spinal arthroplasty designs have been in use for more than 20 years internationally; therefore, concerns about neuropathology, toxicity, and carcinogenicity are mitigated. Debris-induced inflammation still depends on the individual and the type of debris. The consequence of debris-induced inflammation is continued; vigilance by physicians is recommended monitoring of spinal implants using physical exams and testing of metal content and bioreactivity, as is planning for the likelihood of revision in younger individuals.
Bioreactivity (pro-inflammatory reactivity) to orthopedic implants debris is generally the main determinant of clinical performance of established total joint arthroplasties (TJA).1 Aseptic osteolysis has been reported as high as more than 75% of TJA of all implant failures and is the central factor limiting the longevity of current total joint arthroplasties; other reasons for failure include infection (7% of failures), recurrent dislocation (6% of failures), periprosthetic fracture (5% of failures), and surgical error (3% of failures).2 Debris-induced aseptic failure has been reported to be as high as 5% at 7 years postop in current generation metal-on-metal total hip arthroplasties.3, 4, 5 Wear and corrosion products that are attributed with these local inflammatory responses leading to aseptic failure.6, 7, 8, 9, 10 In fact, a 15% revision rate at 20 years has been singularly correlated/predicted by a wear rate of greater than 0.1 mm per year of older model UHMWPE acetabular liners (nonhighly cross-linked).11 To date, wear remains the strongest correlate of aseptic osteolysis and implant loosening. Understanding what the difference is between popular misconceptions and what is known of spinal implant debris is key for physicians to understand the risks associated with modern spine arthroplasty implants. The following review addresses 2 questions: 1) What kinds of spinal implant debris are produced and 2), clinically, what exactly is implant debris reactivity locally and systemically? Unfortunately, the amount of data from total disc arthroplasty studies is limited. Thus extrapolation from hip and knee arthroplasty experience to the spine is still necessary.
Methods included in this review are those associated with physical characterization of implant debris and those associated with measuring the biologic responses to implant debris: 1) gravimetric wear analysis; 2) particle characterization by scanning electron microscopy (SEM) and low-angle laser light scattering (LALLS); 3) metal ion analysis; 4) enzyme linked immunoassay (ELISA); and 5) lymphocyte transformation testing (LTT).
Gravimetric wear is measured by the weight loss of the individual component after simulator or in vivo use. The accuracy of this testing is typically on the order of 0.02 mg. Wear involves the loss of material (weight or mass) as a consequence of relative motion between 2 surfaces. The amount of wear depends on predictable factors: the amount of force pressing the 2 materials together; ie, the contact force, amount of relative sliding distance, type of bearing materials, and the type and amount of lubrication.
Traditional particle sizing techniques, such as SEM or transmission electron microscopy (TEM) (number-based counting methods), indicated that the majority of the wear (mass loss) from an implant is comprised of particles in the nanometer to submicron range. This understanding stems largely from the relatively low numbers of particles (hundreds to thousands) used in traditional number-based analysis techniques such as SEM. Newer analytical techniques, such as LALLS which samples millions to billions of particles, catch the one-in-a-million large particle that comprises a significant portion of the total mass loss (total debris). This contribution of large particles to the total debris mass, but insignificant contribution to the total number of particles, is made apparent in graphical representation of number and volume particle size distributions (Figure 1). The counting of millions to billions of particles facilitates not only a number-based analysis, but because the one-in-a-million larger particle is also counted a volume-based analysis is also facilitated.
These analyses of (a) volume and (b) number distributions of two debris samples demonstrate how similar number distributions can result from very different actual size distributions as evident in (a) the volume distributions. Note: The x-axis represents increasing particle diameter and the y-axis is (a) percentage of total number of particles in each size range and (b) the percentage of total mass of each sample that is of that size range. (Courtesy of BioEngineering Solutions Inc.)
The ability to accurately, and as comprehensively as possible, characterize implant debris is critically important to the new generation of bearing surfaces used in novel spinal implants where weight loss from the implant after a year of use (eg, less than 0.2-mm3 volume loss after a million cycles of use) could be attributed to the loss of a relatively few large particles or hundreds of millions of small particles (200 particles of 100-μm diameter are equivalent in mass to 200 million particles only 1 μm in diameter). This phenomenon is illustrated in Fig. 1, where different samples of particles look demonstrably very different when viewed as a volume-based distribution, though they look very similar when analyzed using a number-based distribution. This shows how important it can be to have number- and volume-based distributions of the same particles to characterize fully the types of particles in the mix.
Growth medium for the human cell lines such as macrophage THP-1 (ATCC) is typically RPMI 1640 supplemented with L-Glutamine, Penicillin, Streptomycin, 25 mM Hepes (Lonza, Walkersville, MD) and 10% heat inactivated fetal bovine serum (FBS; Hyclone Laboratories, Logan, UT). Implant debris that are usually produced using implant components in a cryo-milling or cryo-pulverization techniques (Bioengineering Solutions Inc, Chicago, IL) were generated from metal orthopedic implant components (Cobalt alloy, ASTM F-75). Collected metal particles are characterized for size and number by laser diffraction technology and have a mean diameter of less than 10 μm, usually approximately 1–2 μm (number-based analysis) with a size range of 1–10 μm.
Peripheral blood mononuclear cells (PBMCs) are typically isolated from heparanized whole blood from healthy donors by Ficol-Hypaque density gradient separation and collected for further cell isolation. Peripheral blood CD14+ monocytes are isolated from collected PBMCs by negative selection with a magnetic bead antibody cocktail specific for CD3, CD7, CD16, CD19, CD56, CD123, and Glycophorin A (Miltenyi Biotec). All isolated cell populations are typically assessed for purity by flow cytometry analysis.
Cytokines and chemokines secretion by cells, such as THP-1 monocytes, human primary macrophages, and CD4+ T cells, are analyzed with ELISA or Luminex suspension multiplex array technology. Supernatants from both metal-challenged monocytes and co-culture experiments are collected after 30 minutes to 5 days and stored at −80°C. Monocyte-macrophage supernatants are typically analyzed for IL-6, IL-1β, TNFα, GMCSF, PGE2, and IL-8 with pro-inflammatory multi-plex cytokine array kits (millipore).
Testing for metal sensitivity has historically been conducted in vivo by skin testing (ie, so-called patch testing or intradermal testing) and in vitro by lymphocyte transformation testing (LTT). While general patch testing protocols and commercial kits do exist for a variety of commonly antigenic substances12, 13 (eg, TrueTest, Glaxo Dermatology, Research Triangle Park, NC), there is continuing concern about the applicability of skin testing to the study of immune responses to implants regarding the questionable equivalence of dermal Langerhans cells to peri-implant antigen presenting cells14, 15 and/or the possible induction of hypersensitivity in a previously nonsensitive patient.16
Metal-LTT testing is lymphocyte proliferation testing (also known as lymphocyte transformation testing, hence LTT), and involves measuring the proliferative response of lymphocytes obtained from peripheral blood by routine blood draw (eg, Orthopedic Analysis Inc, Chicago, IL). Metal-LTT assays are typically performed using 96-well cell-culture plates (Sigma) at a density of approximately 0.2 × 106 cells/well for 6 days in 150 μl/well of complete media (DMEM 10% FBS) at 37°C and 0.5% CO2, with or without metal treatments. Radio-labeled thymidine (1 mCi/well) is added during the last 12 hours of the 6-day culture period, where it is incorporated into dividing cells. Cell proliferation is normalized to that of cells treated with media alone (controls) producing a ratio, referred to as the proliferation index or stimulation index (SI). The formula is SI = (mean cpm cells + treatment) / (mean cpm cells + medium alone). All tests should be conducted in triplicate to facilitate an average and standard deviation.
The following results are broken into 2 main sections: a review of the physical aspects of spinal implant debris production and local and systemic biologic responses to implant debris.
The fiction is that there are many different kinds of implant degradation processes. The fact is that there are only 2. Immediately upon implantation all implants begin to degrade through 2 mechanisms, wear and corrosion. Wear is a mechanical/physical degradation of materials (abrasive, adhesive, etc), whereas corrosion is a chemical (electrochemical) form of degradation. Of these 2 degradation mechanisms, wear is the primary source of articulating implant debris.
The fiction is that wear-resistant newer materials, like highly cross-linked polyethylene and ceramic, do not wear and release debris. The fact is that all bearing surfaces wear to some degree. Different types of orthopedic materials and couples produce different amounts and kinds of wear debris. Hard-on-hard material couples, such as metal-on-metal articulations, generally produce less wear (weight loss) than metal-on-polymer (Fig. 2). These 2 basic types of articulating bearing designs are available in spine, knee, and hip arthroplasties.
A graphical comparison of data showing the amount of wear debris generated from different types of total joint arthplasties demonstrating that there is relatively less (10×) polymeric debris generated by a total disc arthroplasty with a metal-on-polymer articulation. However, this difference is not apparent with metal-on-metal articulating implants. Note: Figure References: Metal-Poly:24Ceramic-Poly:25, Metal-X-linked Poly: 22 Metal-X-linked Poly: 26, Metal-X-linked Poly: 27, Metal-X-linked Poly: 28, Metal-Metal: 29, Ceramic-Ceramic: 30, Metal-UHMWPE TDA: 17, Metal-Metal TDA: 18.
In vitro analysis of wear has demonstrated wear rates of metal-on-polymer bearing lumbar total disc arthroplasty devices range from 2 to 20.8 mm3/million cycles,17, 18 where the size of the wear debris generally ranges from 0.1 to 100 μm in diameter UHMWPE.17, 19, 20, 21 This amount of debris is 10-fold less wear than THAs and TKAs that are composed of metal on highly-crosslinked polyethylene (x-UHMWPE) bearing surfaces.22, 23
In general, the wear of metal-on-metal TJA is well below that of metal-on-polymer (Fig. 2).22, 24, 25, 26, 27, 28, 29, 30 The few published reports on the wear rates of metal-on-metal disc arthroplasty prostheses indicate a wear rate of 0.93–1.26 mm3/million cycles (Cobalt-alloy).18, 31 Another study of an all titanium-6% Al-4%V alloy disc arthroplasty found wear rates to be as high as 3 mm3/million cycles.32 These values are similar to those reported for metal-on-metal hip replacements, which have been shown to range from approximately 0.05–6 mm3/million cycles (Cobalt-alloy).29, 33, 34 Long-term follow-up of patients undergoing total disc arthroplasty is required to assess how intimately wear will correlate with inflammation and poor implant performance.
The fiction is that some metals do not corrode in vivo. The fact is that all metals corrode in vivo. Corrosion can be defined as an electrochemical attack of a metal by its environment. Implant corrosion reduces structural integrity and causes the release of by-products that interact locally and systemically. Orthopedic alloys rely on the formation of passive films on their surface for corrosion protection. Some alloys are better than others at resisting this chemical degradation; for instance, stainless steel alloys generally corrode to a greater extent than either cobalt or titanium alloys.35, 36, 37 Special types of corrosion, such as fretting, can take place, where a combination of chemical degradation is enhanced by mechanical factors such as a crevice and abrasive wear, as is common in the screw-plate junction of spine implants. In modern practice, gross evidence of corrosion in screw-plate interfaces is less common.38 Corrosion products are typically oxides (Cr2O3, CoO, TiO2, Al2O3, etc), metal phosphates, metal salts, metal-ions bound to proteins, or organometallic complexes.39, 40, 41, 42, 43
The degradation products of all orthopedic implants are 1 of 2 basic types of debris: particles or soluble debris (metal ions). While there is typically a large distinction between the 2, this distinction becomes a gray area as the size of particles decreases into the nanometer range. Typically, particulate wear debris (metal, ceramic or polymer) exist from the submicron size up to thousands of microns (mm), while so-called “soluble debris” are limited to metal ions that are bound to plasma proteins.
The fiction is that all implant debris are approximately 1 μm in size. The fact is that the easiest (most numerous) debris to measure is typically less than 1 μm in size. But all implant debris cover a range of sizes from nanometers to tenths of a millimeter (0.01 to hundreds of μm in diameter). Currently, there is a lack of comprehensive information regarding the debris retrieved from spinal implants. However, we do know that particles generated in simulator studies of articulating spinal implants match the sizes and types of particles produced from hip and knee arthroplasty (Fig. 3).22, 24, 25, 26, 27, 28, 29, 30 Generally, hard-on-hard material couples, such as metal-on-metal articulations, generally produce smaller-sized (nanometer and sub-micron), fairly round debris, whereas traditional metal-on-polymer bearings produce larger (micron-sized) debris that are more elongated in shape (Fig. 2, Fig. 3). The particles produced from articulating bearing in any metal-on-polymer bearing implants are dominated by polymer particles with little metallic debris, unless there are other sources of metal release such as corrosion at metal-metal connections. Polymeric particles produced from implants generally fall into the range from 0.23–1 μm (Fig. 3A). Past investigations, primarily of UHMWPE wear debris in peri-implant tissues, have shown that 70–90% of recovered particulates were submicron, with the mean size being approximately 0.2–1 μm.43, 44, 45 Newer polymer implant debris from highly cross-linked polymers have demonstrated the production of smaller, more rounded debris in the submicron range as small as 0.1 μm in size.29;46 Metal and ceramic particles have generally been characterized as an order of magnitude smaller than polymer particles (at approximately 0.05 μm in diameter, in the nanometer range). Given the relative wear rates in Fig. 2 and the average particle sizes in Fig. 3A, this translates to a theoretical increase in the yearly production of metal-on-metal particles of 1-3 orders of magnitude over that produced by metal-on-polymer articulating surfaces, even though the amount of mass loss is over an order of magnitude less for metal-on-metal bearings (Fig. 3B). Different types and sizes of particles have been found histologically in peri-implant tissues.19, 35, 47, 48, 49, 50, 51 Stainless steel: At steel screw-plate junctions, which contain particles of chromium compound ranging in size from 0.5 to 5.0 μm, stainless steel has been found in the peri-implant tissues as closely packed, plate-like particle aggregates.52 Cobalt alloy: Cobalt alloy corrosion produces a chromium-phosphate (Cr(PO4)4H2O) hydrate rich material termed “orthophosphate”, which ranges in size from sub-micron to aggregates of particles up to 500 μm.52, 53 Titanium alloy: The degradation products observed in histologic sections of tissues adjacent to titanium base alloys generally have the same elemental composition as the parent alloy (which often is detected as titanium oxide, the thick oxide layer that forms on all titanium implants and protects the bulk alloy), as opposed to precipitated corrosion products which occur with stainless steel and cobalt-chromium alloys (Fig. 4).
A graphical comparison of data showing (a) the reported size of implant debris generated from different types of total joint arthplasties and (b) the amount of particles per year this results in using the gravimetric data shown in Figure 1. The smaller size of the particles reported in highly crosslinked polymer particles (copared to traditional poly) combined with the modest reductions in wear rate results in metal on Metal-X-linked Poly. Note: Figure References: Metal-Poly:24Ceramic-Poly:25, Metal-X-linked Poly: 22 Metal-X-linked Poly: 26, Metal-X-linked Poly: 27, Metal-X-linked Poly: 28, Metal-Metal: 29, Ceramic-Ceramic: 30, Metal-UHMWPE TDA: 17, Metal-Metal TDA: 18.
Transmission Electron Photomicrographs: (a) Macrophage containing phagocytized titanium particles. (b) Endothelial cell lining with embedded titanium debris. These specimens were obtained from a tissue sample overlying the posterolateral fusion mass (sixteen-week autograft + titanium) (TEM magnification = 20,000×) (courtesy of Bryan Cunningham).
The fiction is that scanning electron microscopy is the best way to analyze the size of particles. The fact is that newer techniques, such as low-angle laser light scattering (LALLS), can increase the number of counted and sized particles from hundreds to millions or billions facilitating catching the one in a million large particle. Traditional particle-sizing techniques, such as SEM or TEM (number-based counting methods), indicated that the majority of the wear (mass loss) from an implant is comprised of particles in the nanometer to submicron range. This understanding largely stems from the relatively low numbers of particles (hundreds to thousands) used in traditional number-based analysis techniques such as SEM. Newer analytical techniques catch the one-in-a-million large particle that comprises a significant portion of the total mass loss (total debris). This contribution of large particles to the total debris mass, but insignificant contribution to the total number of particles, is made apparent in graphical representation of number and volume particle size distributions (Fig. 1). The counting of millions to billions of particles facilitates not only a number-based analysis, but because the one-in-a-million larger particle is also counted, a volume-based analysis is facilitated.
The ability to characterize accurately, and as comprehensively as possible, implant debris is critically important to the new generation of bearing surfaces used in novel spinal implants, where weight loss from the implant after a year of use (eg, less than 0.2 mm3 volume loss after a million cycles of use) could be attributed to the loss of a relatively few large particles or hundreds of millions of small particles (200 particles of 100 μm diameter is equivalent mass to 200 million particles only 1μm in diameter). This phenomenon is illustrated in Figure 1, where different samples of particles look demonstrably very different when viewed as a volume-based distribution, though they look very similar when analyzed using a number-based distribution. This shows how important it can be to have number- and volume-based distributions of the same particles to fully characterize the types of particles in the mix.
The fiction is that due to the smaller loads, spine arthroplasty implants release less metal ion into the body than hip or knee arthroplasties. The fact is that metal levels measured in people with disc arthroplasties have comparable levels of circulating metal ions as people with other TJA implants. There is continuing concern regarding the release of chemically active metal ions, which bind to proteins and remain in solution from which they can then disseminate into the surrounding tissues, bloodstream, and remote organs. Particulate metallic wear debris present an enormous surface area for electrochemical dissolution, which, in all likelihood, is a major factor contributing to observed systemic elevations in metals of patients with titanium implants.35 Normal human serum levels of prominent implant metals are approximately: 1-10 ng/ml Al, 0.15 ng/ml Cr, <0.01 ng/ml V, 0.1-0.2 ng/ml Co, and <4.1 ng/ml Ti. Following total joint arthroplasty, levels of circulating metal have been shown to increase (Table 1). The values in this table show that following successful primary total joint replacement there are measurable elevations in serum and urine Co, Cr, and Ti. Transient elevations of urine and serum Ni have also been noted immediately following surgery.
|Concentrations of metal in body fluids (ng/mL or ppb)|
|Normal||<0.2–0.6109 (0.2–8.3)109||0.2110 (0.1–0.7)110||0.42110 (0.3–1.8)110||1.2–2.7109, 111 (1.1–7.9)||1.2–2.2111 (1.1–6.4)||<0.8111 <0.8||109, 111|
|TKA-M/P||0.2 (<0.02–1.15)112||0.1 (<0.02–0.6)112||0.4 (0.05–1.5)112||3.2 (<2.1–6.3)111||6.5 (2.1–9.4)111||<0.8 <0.8111||111, 112|
|TKA-F||135.6 (24.1–716.9)111||3.7 (0.8–6.2)111||0.9 (<0.8–2.6)111||111|
|THA-M/P||0.9 (<0.3–3.9)113||1.28 (0.1–2.4)113||1.4–4.16, 2.6113 (<1.1–11.17)109, 113||1.2–1.7109 (0.2–2.46)109||40, 109, 113, 114|
|THA-M-M||2.4 (0.6–7.9)113||3.5 (0.8–9.1)113||1.9113 (1.1–6.0)113||113|
|TDA||1.9–4.854, 55||1.9–2.454, 55||54, 55|
|Normal||<0.1 (<0.1–7.6)115||7.3 (1.9–19)115||115, 116|
|THA||5116, 0.2 (<0.02–1.15)112||3116, 0.1 (<0.02–0.6)112||21 ± 8116||4116, 0.4 (0.05–1.5)112||13–556116||109–654116||5–62116||112, 116|
|THA-F||588 ± 427116||385 ± 232116||58 ± 53116||32 ± 16116||86 ± 35116||256 ± 271116||25 ± 19116||116|
|Normal||0.1–1.2116||2–6116||0.5–1.8116||2.9–7.0116||17 ± 60116||12 ± 4116||5.8 ± 4116||116|
|THA||67 ± 62116||218 ± 233116||23 ± 31116||116|
|THA-F||20 ± 25116||110 ± 150116||12 ± 9116||29 ± 29116||602 ± 927116||237 ± 307116||55 ± 63116||116|
|Human tissue (μg/g) (roughly equivalent to: 10–1 to 10–2 mM)|
|Normal||<65||120||<9||50||150||116, 117, 119, 120|
|TJA||39400||460||121||5490||3820||116, 117, 119, 120|
Abbreviations: Normal, subjects without any metallic prosthesis (not including dental); M/P, metal on polyarticulation; M/M, metal-on-metal articulation; F, subjects with a poorly functioning arthroplasty (loose or migratory osteolysis prior to surgical revision); THA, total hip arthroplasty; TKA, total knee arthroplasty; TJA, subjects with well functioning total joint arthroplasty (either knee or hip); TDA Spine, total disc arthroplasty (metal on metal).
Recent studies of metal-on-metal TDA find serum levels of cobalt and chromium after TDA are elevated at concentrations of 3–4 ng/mL or parts per billion (ppb) for cobalt and 1–2 ng/mL or ppb for chromium, which were significantly elevated over control values.54, 55 The concentrations of circulating Cr/Co measured in the serum of people with a metal-on-metal TDA are similar to levels measured in well functioning metal-on-metal total hip arthroplasties. This is not true for nonarticulating implants, where recent studies have failed to detect elevated amounts of circulating Cr or Ni from stainless steel scoliosis rod fixation systems.56 The long-term effects of presumably much higher elevations of metal proximal to the spinal implant remain unknown and are under careful surveillance by implant companies, the FDA and orthopedic researchers.57
There are subject dependent conditions that complicate using metal ion analysis as a diagnostic technique; eg, altered renal function has been reported to cause variability in circulating metal levels. Brodner et al have reported that renal failure is associated with highly elevated levels of serum Co and Cr (over 100-fold) in comparison with individuals with similar prostheses and normal renal function.58 At this time, there is no established toxicity threshold for the degradation products of metal implants.
The fiction is that, generally, different types of implant debris elicit the same level of inflammation. The fact is that different types of implant debris elicit different types of biologic response. However, despite the large number of published studies examining the effect of metal, polymer and ceramic particles on peri-implant cells (eg, macrophages, fibroblasts, osteoblasts, and osteoclasts), there is surprisingly little consensus on what type of debris are most pro-inflammatory. That said, there are a few general particle characteristics on which local inflammation has been shown to depend: 1) particle load (particle size and total volume), 2) aspect ratio, and 3) chemical reactivity. Thus, theoretically, a bioreactivity index of particulate would take the mathematical form:
where Ks are constants that depend on the testing environment and particle load is a function of both particle size and total debris volume.
The inflammatory response is proportional to the particle load (the concentration of phagocytosable particles per tissue volume, which is characterized by both the size and total volume).59, 60 If a given amount of debris (mass loss from an implant) is comprised of small diameter particles, there will be far greater numbers than if that same mass of debris was composed of larger diameter particles.60 The degree to which equal numbers (dose) of large versus small particles (eg, 10 μm vs 1 μm diameter) induce a inflammatory response on a per-particle basis in vivo has not been thoroughly investigated and remains an issue of debate. However, several past investigations have shown that for an equal amount of debris mass, smaller particles (eg, 0.4 μm) produce a greater inflammatory response than larger particles (eg, 7.5 μm).60, 61, 62 In these studies, the challenge doses used generally contain many thousands more small particles per cell than larger particles; this is largely responsible for these responses and the popular misconception that smaller particles are always more pro-inflammatory regardless of dose.
This phenomenon has been well established with the first such investigations over 30 years ago, involving asbestos fibers, which similarly cause inflammation through “danger signaling,” inflammasome activation and production of IL-1beta, IL-18, IL-33, etc, and activation of general macrophage pathogen associated molecular pattern pathways (NFkb and p38 production of pro-inflammatory cytokines IL-6, TNFα and PGE2 through as yet unknown mechanism).65, 66 This is discussed further in the following section on local implant induced inflammation. However, it remains unknown at what point (aspect ratio: length/width) in the transition of round particles to fibers that elevated inflammation is generally initiated. Currently, fibers can be categorized as particles with aspect ratio greater than 3, as has been shown in metal and poly retrieved from tissues.29, 44, 67
There is a growing evidence that metal particles are more proinflammatory to macrophages (in general) when compared to polymers (ceteris paribas) and that some metals are more reactive than other metal particles.68 This opinion is not unanimous.69 More complicated still, there have been reports that have shown no differences between similar sized metal and polymer particles.70, 71, 72 Despite these reported differences, there is a growing consensus that metallic particles that are capable of corroding and releasing ions that have been associated with hypersensitivity responses, cytotoxicity, and DNA damage are thus more capable of eliciting proinflammatory responses when compared to relatively inert polymers and ceramics.73, 74, 75
Does particle size matter? While there seems to be a large consensus that in vitro inflammatory responses require particles less than 10 μm in diameter, there continues to be debate on exactly what size particle produces the greatest inflammatory response.60, 61, 62, 76 Thus, to produce an in vitro inflammatory response, particles need to be less than 10 μm, that is, within a phagocytosable range. Purportedly, particles with a mean size of 0.2–10μm are generally the most proinflammatory. Within this range there is no consensus as to which specific sizes and/or dose of particles (particles/cell or particles/tissue volume) are maximally inflammatory.60, 77, 78, 79
Endotoxin (bacterial-derived lipopolysaccharides) on wear particles is presumably important because it has been found in periprosthetic tissue of failed implants, even in the absence of clinical signs of infection.80 The relationship between bacteria and aseptic loosening has also been inferred, because antibiotic-eluting bone cement and systemically administered antibiotics reportedly reduce the frequency of aseptic loosening.81
The fiction is that implants generally “wear out” over time. The fact is that describing to patients that implants simply wear out over the long term mischaracterizes the main reason for long-term revision, aseptic loosening. Implant debris from wear cause local inflammation and granulomatous invasion of bone-implant contact that, over time, results in implant loosening and pain, necessitating revision arthroplasty.
Implant debris is known to cause inflammation, osteolysis, and, in some cases, hypersensitivity, and concerns about implant debris becoming carcinogenic and/or toxic persist. Macrophages are part of the innate immune system and, in the context of implant debris, are dedicated to phagocytosing debris, eliciting danger signals, and presenting antigen for T-cell recognition (ie, metal-protein complexes). Once debris is ingested by macrophages and other peri-implant cells, a host of pro-inflammatory reactions occur such as activation of metal-reactive T-cells through antigen presentation,73 release of pro-inflammatory mediators,60, 82, 83, 84 cytotoxicity,74 DNA damage,74, 85, 86 and oxidative stress.87 The central pro-inflammatory cells in this response are macrophages, which are generally responsible for mediating debris-induced inflammation leading to device loosening.
The following section will detail what is known of these different pathologies induced by implant debris and the primary suspects (kinds of implant debris for each pathology). Given that TJAs have been in use for over half a century with constant increases in biocompatibility and materials technology over that time frame, implant concerns about toxicity and carcinogenicity have been largely mitigated.
The fiction is that the amount of wear and corrosion debris produced from spine arthroplasty and modular spinal fixation implants are inconsequential and not likely to provoke a deleterious immune response. The fact is that biologic reactivity to spinal implant debris has been clinically observed with all the hallmarks of traditional particle-induced osteolysis; ie, granulomatous epithelioid membranes coating the metal implants have been reported, similar to the fibrous membranes associated with loose total hip replacements.19, 49, 50, 88, 89, 90, 91, 92
Although few, case reports of painful granuloma associated with spinal implant debris93, 94 demonstrate that spinal implant debris induced inflammation can result in bone-destroying granuloma.95 It is generally accepted that decreased wear and corrosion debris will lead to increased longevity of orthopedic implants. This rule of thumb has yet to be well established over the long term in newer types of articulating spinal implants; however, it is well established that debris induced inflammation and pain can be caused by spine implant debris.49, 94, 96, 97, 98, 99
Macrophages are the major cell type responsible for mediating the pro-inflammatory response to implant debris that leads to inflammation and device loosening over time. Once implant debris is ingested by macrophages and other peri-implant cells, a host of biologic reactions can occur such as activation of T-cells through antigen presentation,73 release of pro-inflammatory mediators,60, 82, 83, 84, 100 cytotoxicity,74 DNA damage,74, 85, 86 and oxidative stress.87 Of these reactivity types, it is the pro-inflammatory response to wear debris in the local environment that has been the only known correlate of implant performance.11
In the relatively few reports of human retrieval studies of loose spinal implants, granulomatous epithelioid membranes coating the metal implants have been shown similar to the fibrous membranes associated with loose total hip replacements.19, 49, 50, 89, 90, 91, 92 However, before this phenomenon of wear debris can be well established in disc arthroplasty, larger series of implant retrievals will be needed to characterize the biologic responses to increased particle generation. When osteolysis or granulomas occur in the presence of metal debris, it is termed particle disease. In the case of metal debris, this is often accompanied by metallosis (aseptic black staining). Metallosis often accompanies metal implant debris-related osteolysis, aseptic fibrosis, local necrosis, or loosening. In a cohort of 12 loosened spinal implant cases, metallosis of the internal membrane was associated with the outer layer of membrane containing an infiltrate of leukocytes and macrophages (see Fig. 5). All 12 patients had radiolucency around part of the spinal instrumentation. Seven of the patients had titanium implants and 5 of the patients had stainless steel implants, showing this phenomenon is not limited to 1 type of material. Eleven of 12 clinical cases demonstrated elevated TNFαlevels and an increased osteoclastic response in the vicinity of wear debris caused by dry frictional wear particles of titanium or stainless steel (see Fig. 6). The focal areas of osteolysis most commonly involved loose transverse connectors. Resection of the wear debris stained tissue and surrounding fibro-inflammatory zones resulted in resolution of clinical symptoms in all 12 cases.88