Mechanical Properties of Implant in Lumbar Interbody Fusion: History
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Contributor:
Jakub Litak
,
Wojciech Czyzewski
,
Michał Szymoniuk
,
Zofia Hoffman
,
,
Leon Sakwa
,
Piotr Kamieniak

Over the last decade, pedicle fixation systems have evolved and modifications in spinal fusion techniques have been developed to increase fusion rates and improve clinical outcomes after lumbar interbody fusion (LIF). Regarding materials used for screw and rod manufacturing, metals, especially titanium alloys, are the most popular resources. In the case of pedicle screws, that biomaterial can be also doped with hydroxyapatite, CaP, ECM, or tantalum. Other materials used for rod fabrication include cobalt–chromium alloys and nitinol (nickel–titanium alloy). In terms of mechanical properties, the ideal implant used in LIF should have high tensile and fatigue strength, Young’s modulus similar to that of the bone, and should be 100% resistant to corrosion to avoid mechanical failures.

  • metal alloys
  • implants
  • inter body fusion

1. Introduction

Over the past few decades, lumbar spinal fusion (lumbar interbody fusion, LIF) has been recommended as a well-known, standard surgical treatment for degenerative disc disease (DDD) of the lumbar spine. DDD may cause low back pain and radicular symptoms, which can significantly decrease the quality of life. The prevalence of symptomatic DDD increases with age and occurs in 10% of the male population at the age of 50 and up to 50% at the age of 70 [1]. According to some reports, DDD may concern even 90% of the population including asymptomatic cases [2]. LIF effectively provides stabilization of painful motion segment, restores lordosis and disc height, corrects the deformity, and may provide indirect decompression of dural sac and nerve roots [3][4]. That allows immediate relief of DDD symptoms. Other indications for this procedure include traumatic injuries, degenerative or congenital deformities, spondylolisthesis, spinal stenosis, and tumors [3][5][6][7].
There are various approaches to a lumbar interbody fusion. However, there is a lack of sufficient and reliable evidence to establish one of them as a standard lumbar fusion method. Posterior lumbar interbody fusion (PLIF) and anterior lumbar interbody fusion (ALIF) are the most traditional techniques. Nowadays there are other, less invasive methods including lateral lumbar interbody fusion (LLIF), extreme lateral lumbar interbody fusion (XLIF), oblique lumbar interbody fusion (OLIF), and transforaminal lumbar interbody fusion (TLIF) [6]. Moreover, minimally invasive approaches such as minimally invasive TLIF or percutaneous pedicle screw fixation have gained popularity recently [8]. All lumbar spinal approaches require the use of proper instrumentation. The basic spinal fixation device consists of pedicle screws, connection rods, a cross-link device, and in some cases an interbody cage. Pedicle screws are placed into the vertebral bodies through the pedicles of vertebrae, the Harrington rods connect screws of adjacent vertebrae, and the cage is inserted into the intervertebral space. Such an interbody device enables distraction of disc space and successfully stabilizes the pathological segment.

2. Physical and Mechanical Properties of Implant Important in LIF

The ideal biomaterial used in LIF should have high tensile and fatigue strength, Young’s modulus similar to that of the bone, and should be 100% resistant to corrosion to avoid mechanical failures. Therefore, the composition of the spinal rods and screws constitutes a crucial factor in defining the general functionality of the spinal instrumentation.

2.1. Fatigue Strength

One of the most important features of LIF implants is their fatigue strength. That property describes how long the spinal instrumentation can work without breaking down [9]. The cycling loading of the spine, which appears during daily activities, generates oscillating stresses on spinal instrumentation and may lead to a crack in the implant material. When the crack reaches a critical size, fatigue fracture of the material occurs, leading to the failure of the implant [9][10]. Remarkably, long cracks cause implant collapse slower than very small cracks, which are relative to the dimensions of the material micro-architecture [11]. Biomechanical performance and fatigue strength of spinal instrumentation significantly depend on the microstructure of the metal alloy of which it is made [12][13][14]. To increase the fatigue life of an alloy, many heat treatment techniques are implemented [15]. They include plasma-assisted microwave chemical vapor deposition, plasma nitriding, plasma etching, and deposition of amorphous diamond-like carbon (a-DLC) layers inoculated with nitrogen and silicon. It has been shown that these methods have a substantial influence on the surface characteristics and microarchitecture of alloys [16]. Furthermore, heat treatment allows for arranging an optimized balance of material features such as machinability, ductility, and stability [15].
Fatigue fractures almost always occur at stress concentration sites such as notches or the discontinuity of the geometrical structure of the material [10][11][17][18]. On the spinal rod’s surface, notches may be generated during manual contouring, which can impair the mechanical properties of the rod, especially its fatigue strength [10][19][20]. This phenomenon is known as the “notch effect”. Before metal rods are fixed to the patient’s spine, they are contoured to obtain optimal sagittal alignment of the spine [10][21]. Traditionally, contouring can be done by the surgeon with the use of the French bender. Alternatively, the rods can be anatomically designed and bent automatically by a machine before the surgical procedure. A biomechanical study by Yamada et al. [22] has shown that pre-contoured rods had a remarkably higher fatigue strength and ultimate load than intraoperative manually contoured rods at the same load condition. Moreover, pre-contoured rods not only reduce the risk of rod fracture but may also reduce operation time and bleeding and decrease the risk of infection in comparison to the manually bent rods [22][23]. This results from the manual contouring of the rods that requires a significant amount of time to adjust the proper shape of the rod. Furthermore, tightening of the screw leaves surface defects on the rods and may contribute to the notch effect formation, which has been described by many researchers[10][11][12]. Therefore, avoiding severe tightening of the screws is recommended.
Recently, some studies have shown the negative influence of direct electrocautery use on the mechanical features of metal alloys [24][25][26]. According to a biomechanical study by Zobel et al. [25], electrocautery contact with the material was found to significantly decrease the fatigue strength of the Ti-6Al-4V titanium alloy. Even after a short contact with the electrode, the fatigue strength reduced remarkably. When electrocautery contact is applied at high-stress concentrations areas of instrumentation, it can be a notable problem for the mechanical properties of the implant and in extreme cases, it may lead to implant breakage [25][26]. Moreover, this problem was reported by many case studies in hip replacement surgery [26][27][28]. However, it remains unclear whether a decrease in fatigue strength after electrocautery contact depends on the material and whether it is determined by the type of the implant [25]. Furthermore, there is a lack of reports describing that issue in the case of other metal alloys. Regardless of that, spine surgeons should pay special attention and avoid any contact of the active electrocautery electrode with implants during the revision surgery, especially in the case of titanium implants and areas of the implant with high-stress concentrations.

2.2. Young’s Elastic Modulus

A further property, which is crucial for the metal alloy to be useful as the material for the spinal implant, is its elasticity, which is the ability of a metal to resist distorting influence and return to its initial shape. This feature of a material can be described by a physical quantity known as Young’s elastic modulus. The value of the elastic modulus of human cortical bone ranges from 10 to 30 GPa [29]. The perfect alloy for use in spinal fusion systems should have Young’s modulus as similar as possible to the bone. That prevents a phenomenon called a “stress shielding effect”. That term refers to the reduction in bone density around the implant due to bearing of the majority of the mechanical load by the instrumentation. Normally, the bone experience stresses and remodels in response to the loadings. Therefore, due to decreased load, bone atrophy progressively occurs and it may result in implant loosening and failure [30].
Contemporary material engineering enables the development of metallic biomaterials, which have a modulus of elasticity more similar to human bone. They include different compositions of metals in alloys, which impact their mechanical properties and also materials with various porosity. Creating pores in the alloys not only improves osteointegration of the implant due to the better ingrowth of bone tissue into them but can also affect the value of Young’s modulus of the material. When the porosity of alloy increases, strength and elastic modulus of alloy decrease linearly [31][32]. Therefore, designing and manufacturing implants with the value of the elastic modulus close to that of human bone is possible.

2.3. Corrosion Resistance

The perfect biomaterial should also be corrosion resistant in any environment, especially in the internal environment of the body over any period of time. Generally, corrosion is a progressive degradation of a material resulting from its interaction with the extracellular body environment [33]. That environment contains a lot of ions of sodium, calcium, potassium, magnesium, chloride, phosphate, and bicarbonate, which may be potentially very corrosive factors [9][33].
In the literature, there are three different types of corrosion—fretting corrosion, crevice corrosion, and galvanic corrosion. Each of them has been observed in metallic spinal fusion systems [34][35]. Fretting corrosion develops as a result of mechanical damage from repeated micromotion and friction over time, occurring during the patient’s daily activity. It leads to the release of debris into the surrounding tissue [34]. This type of corrosion is determined by multiple factors such as the design of spinal instrumentation, used metal alloy, electrochemical environment, and load conditions [36]. Crevice corrosion results from exposing the metal to a surrounding tissue fluid, which can induce a local corrosion process by the point destruction of the passive oxide film [37]. Galvanic corrosion is the result of the presence of two different metals in contact with each other in the fluid environment [34]. The use of dissimilar metal alloys in the same spinal instrumentation systems could improve its mechanical features. On the other hand, mixing dissimilar metals in spinal implants brings with it an increased risk of inducing galvanic corrosion [34][36][38]. However, biomechanical studies conducted in 0.9% sodium chloride at 37 °C and retrieval analyses of spinal instrumentation have shown no evidence of galvanic corrosion in spinal constructs made of different metals [36][39][40]. Researchers compared breakdown potential to assess corrosion resistance of each discussed metal alloy. Materials with breakdown potential below 300 mV are regarded as unacceptable. The value of breakdown potential above 600 mV is considered corrosion resistance. Materials with marginal breakdown potential which ranges from 300 mV to 600 mV should be tested under the indicated use [41].
Most metallic alloys used in LIF are passive metals, which means that they have a stable oxide film on their surface [42]. That layer plays an important role in corrosion protection and the loss of its stability results in inducing the corrosion. In the presence of the above-mentioned ions in the surrounding environment, especially chloride ions, the passive film may be damaged [43]. Mean chloride ion concentration in interstitial body fluids is 113 mEq/L, which can induce corrosion in metallic implants [44]. Moreover, cycling loading, micromotion resulting from fretting, and other mechanical factors may also discontinue the passive layer on the surface of the implant [45].
Corrosion has a negative impact and not only leads to failure of the implant but also may leach debris and metal ions that could be harmful to the surrounding tissue. Moreover, some studies of spinal implants have detected elevated serum metal ion levels [37][46][47]. Other studies have found metal debris in lymph and organs such as the liver, spleen, and kidneys [48][49]. Metal ion release induces biological complications such as toxicity, hypersensitivity, and also cancerogenicity [33]. This phenomenon has been correlated with the output of cytokines and metallic proteases by activated macrophages, neutrophils, and T lymphocytes [50]. Other noted complications include pseudotumor and particle-induced osteolysis [51][52]. Localized neurological damage associated with rod breakdown has also been noted in several case reports [53][54][55].

3. Mechanical Characteristics of the Most Frequently Used Metal Alloys in LIF

3.1. Titanium

Among all the metallic alloys used in the manufacturing of spinal instrumentation for LIF, titanium alloys are the most common materials. They owe their popularity to their excellent biocompatibility, superior mechanical properties, great corrosion resistance, and appropriately low Young’s modulus and generate minimal artefacts on computed tomography or magnetic resonance imaging [11][31][56][57]. These properties are highly preferable for biomedical applications. Due to the low elastic moduli and quite often observed notch effect of the titanium rods, titanium alloys are more often used in spinal screw fabrication than the spinal rods [10][58]. In contact with the air, a passive oxide film (TiO2) forms on the surface of the titanium. This layer is probably responsible for resistance to corrosion, chemical inertness, and stability of that metal [31].
There exist two well-known allotrophic phases of titanium—α and β phases. The type of alloy depends on the allotrophic phase of titanium which has been applied. Thus, researchers distinguish between α, near-α, α–β, and β alloys of titanium.
Ti-6Al-4V alloy (α–β type alloy) is the most frequently used titanium alloy for spinal fixation devices [42][59][60]. Despite biocompatibility, excellent corrosion, and mechanical resistance of that alloy, its elastic modulus (~110 GPa), which is higher in comparison to human bone, may induce a stress-shielding effect and result in pedicle screw loosening and bone absorption [9][60][61][62]. Moreover, some studies have shown that Ti-6Al-4V accelerates the development of adjacent segment disease [62]. However, compared with other non-titanium metallic alloys used in spine fusion systems, Ti-6Al-4V has a relatively low modulus of elasticity and the stress-shielding effect is not as strong. Coating the pedicle screws with various materials such as PMMA, hydroxyapatite, extracellular matrix, and titanium plasma spray in tantalum was developed to improve the fixation and pull-out strength of the Ti-6Al-4V screws [63]. Many studies have successfully shown that coated screws may significantly increase resistance against pull-out force in comparison to uncoated screws [64][65][66]. Regarding potential toxicity associated with the leaching of the vanadium and aluminum ions from the Ti-6Al-4V implants, the amount of these metals released is minimal and does not induce suspected health problems, such as neurological or enzymatic disorders [67].
To better adjust the elastic modulus of titanium biomaterials to the cortical bone, increasing the β phase percentage in the alloy is an effective way [56]. One of them, in which this method was applied, is Ti-24Nb-4Zr-8Sn (Ti2448 alloy, α–β type). This material, drawing the attention of many researchers, has a lower Young’s modulus (~49 GPa) than Ti-6Al-4V and shows no toxic features [68]. Therefore, due to the elastic modulus value being more similar to human bone, the stress-shielding effect may be significantly less observed. It has been confirmed in a study conducted by Qu et al. [69] on a porcine model, which compared stress-shielding effects between Ti-24Nb-4Zr-8Sn alloy and Ti-6Al-4V alloy. Another low-modulus titanium alloy is the Ti-45Nb alloy (β type alloy) [70][71][72]. Besides the decreased value of Young’s elastic modulus, Ti-45Nb presents beneficial osteogenic features, which result from a high content of niobium [73]. Additionally, titanium alloys with increased content of β phase show increased corrosion resistance [74]. On the other hand, titanium alloys with low elastic modulus, such as β type alloys, usually also have low mechanical strength [62]. One of the well-known effective methods to increase the mechanical resistance of metals is precipitation hardening. However, due to irreversible changes in the crystalline structure, Young’s modulus increases, and corrosion resistance reduces due to this technique [75]. On the other hand, many studies have shown that severe plastic deformation (SPD) techniques, which include high-pressure torsion (HPT) [71][76][77][78], hydrostatic extrusion (HE) [79], and rolling and folding (R&F) [72], may improve the strength of β type titanium alloys without changing Young’s modulus. SPD techniques also improve corrosion resistance through the thickening of passive film [78]. However, alloys containing niobium, molybdenum, wolfram, or tantalum are expensive to manufacture due to the rarity and high melting points of these metals [42]. Thus, β type alloys such as Ti-24Nb-4Zr-8Sn and Ti-45Nb, despite their appropriately low Young’s modulus, excellent corrosion resistance, and sufficient mechanical properties after SPD processing, may have problems with spread of their use in LIF devices due to the very high costs of production.
Decreasing Young’s modulus of titanium alloys may be also achieved by creating pores in them. A biomechanical study by Skolakova et al. [67] has shown that Ti alloy with the addition of 30 wt.% pore-forming agent (PA) obtained with self-propagating high-temperature synthesis (SHS) exhibits a very similar elastic modulus (~9 GPa) to human bone with good corrosion resistance. However, mechanical strength decreased after the SHS procedure and further studies are necessary to evaluate the usefulness of Ti with 30 wt.% PA in LIF.

3.2. Cobalt–Chromium

Another metal alloy which may be used in LIF systems is cobalt–chromium (CoCr) alloy. This biomaterial usually consists of 63% cobalt, 28% chromium, 5% molybdenum, and minor amounts of other metals [42]. Well-known applications of CoCr alloy as biomaterial include hip and knee joint implants, as well as crowns and implant abutments in dentistry [42][67][80]. CoCr is characterized by higher fatigue life and strength, increased stiffness, and better resistance to notch effects in comparison to titanium alloys [12][17][81][82][83][84]. Due to its higher Young’s modulus (~210 GPa) [24] than titanium, CoCr spinal rods more effectively stabilize the spine and correct abnormalities of spinal curvatures such as scoliosis [83][85][86]. In the study by Willson et al. [87], CoCr rods demonstrated the least amount of shape loss in a radius of curvature compared with commercially pure titanium rods throughout the study.
However, high stiffness of CoCr rods may result in acceleration development of adjacent segment disease [86][88]. According to a comparative study by Han et al. [89], breakages of CoCr rods have been less observed than for titanium rods, but in the case of CoCr, they observed a more frequent occurrence of proximal junctional kyphosis (PJK), which is a form of adjacent segment degeneration. Moreover, a Young’s modulus significantly higher than that of human bone disqualifies CoCr alloy as a biomaterial for screw manufacturing due to the increased risk of stress-shielding effects. Furthermore, compared with titanium alloys, CoCr has lower corrosion resistance, which results in higher overall metal ion release from CoCr implants. Leaching cobalt ions from the alloy due to fatigue and biocorrosion may cause metallosis, neurological-related symptoms (such as deafness and blindness), hypothyroidism, cardiological and hematological problems, and also cancers [90][91][92][93]. To prevent these issues, coating of CoCr implants with ceramics such as calcium phosphate can decrease cobalt ion release and improve biocompatibility, which has been proven in a study by Bandyopadhyay et al. [94] with the use of the surface melting (LSM) technique for tribofilm formation.

3.3. Nitinol

Nickel–titanium (nitinol) is a metal alloy which consists of titanium and nickel in equal atomic percentages [42]. Among all the alloys used in spinal fusion devices, nitinol is characterized by a unique feature which is its superelasticity [42][67][95]. This phenomenon enables the nitinol implant to immediately return to an undeformed shape after removal of external load, even after large deformations. In this way, the use of this super-elastic alloy in LIF systems as rod material makes stabilization more dynamic and may prevent ASD occurrence [96]. Moreover, nitinol is used clinically in intravascular stents, osteosynthesis staples, and orthodontic wires [67][96][97].
Additionally, nitinol has Young’s modulus ranging from 40 to 75 GPa, which is optimal for biomedical applications. Moreover, nitinol fabrication by combustion synthesis (CS) enables tailoring its elastic modulus to that of human bone with great accuracy. After that procedure, metal alloy achieves high compressive strength with appropriately low Young’s modulus and excellent corrosion resistance [97]. According to Aihara et al. [97], general porosity of nitinol to obtain the best elasticity was found to be 64%.
Due to the formation of a passive titanium oxide film (TiO2) on the surface of the nitinol, it is considered a long-term corrosion-resistant and biocompatible alloy [96]. Therefore, coupling nitinol rods with titanium pedicle screws may be considered the best combination for spinal fusion devices due to its high resistance to galvanic corrosion [98]. The corrosion resistance of nitinol alloy is better than CoCr and 316L stainless steel, but inferior to that of Ti-6Al-4V [42][96]. However, the in vitro study combined with retrieval analysis of the nitinol, CoCr, and Ti-6Al-4V rods by Lukina et al. [96] has also shown that nitinol fretting corrosion patterns were worse compared with CoCr. That result may be an effect of lower resistance to fretting corrosion of the nitinol due to higher mobility of the rod. Moreover, intensive fretting may damage the passive oxide layer, whose restoration is relatively low. Thus, it may induce galvanic corrosion and deteriorate its overall corrosion resistance. As result, it affects the fatigue strength of nitinol and may release nickel ions into the blood. However, some studies have shown that the nickel ion levels in blood and tissues were not higher compared with the control group. In any case, to prevent fretting corrosion, coating nitinol rods with protective layers and enhancing the locking mechanism of the pedicle screws would be beneficial solutions [96].

3.4. Stainless Steel

Before introducing titanium alloy as a biomaterial for spinal construct manufacturing, stainless steel (SS) was the most popular metal alloy in this field. It is widely used for other biomedical applications such as bone fixation, cardiovascular systems, catheters, surgical instruments, or dental crowns. Surgical 316 L SS is the most common form of stainless steel for biomedical uses. This specific composition consists of 0.02% carbon, 10–14% nickel, 16–18% chromium, 2% manganese, 2–3% molybdenum, with the rest being iron. The high mechanical properties of this alloy are great advantages for use in spinal fixation. However, SS exhibits a significantly higher elastic modulus (210 GPa) [97] in comparison to human bone. Thus, the stress-shielding effect is strongly observed after SS implant application [99]. Regarding corrosion resistance, many studies have shown that it is significantly inferior compared with CoCr and titanium alloys [34][36][39]. Long-term biomechanical tests by Singh et al. [39] have shown that both CoCr and titanium constructs were more resistant to the fretting corrosion compared with SS. Moreover, during the corrosion process, SS constructs have produced a noticeably greater volume of debris than titanium or CoCr instrumentation systems [39]. Therefore, stainless steel should no longer be in used in spinal surgery.

This entry is adapted from the peer-reviewed paper 10.3390/ma15103650

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