Standard Testing

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STANDARD TESTING

13

Annette Kienle⁎, Nicolas Graf⁎, Tomaso Villa†,‡, Luigi La Barbera†,‡

SpineServ GmbH & Co. KG, Ulm, Germany* Politecnico di Milano, Milan, Italy† IRCCS Galeazzi Orthopedic Institute, Milan, Italy‡

I­NTRODUCTION: ADVANTAGES AND DISADVANTAGES OF STANDARD TESTING Implantable medical devices must meet specific testing requirements before they can be put on the market. These requirements are defined by respected regulatory bodies such as the United States Food and Drug Administration (FDA) and the European Notified Bodies. To improve the comparability between testing laboratories, testing is carried out according to national or international standards wherever possible. The main aim of standard testing is to investigate the safety and effectiveness of medical devices. The test results are part of the submission process used in seeking approval from the appropriate regulatory body. In Europe and North America, ASTM (American Society for Testing and Materials) and ISO (International Organization for Standardization) standards need to be followed for the most part. The process of standard development and revision plays an important role in better understanding the advantages and disadvantages of standard testing. In the case of the American Society for Testing and Materials, task groups develop the basics for the draft standards. These drafts pass through different hierarchical levels (subcommittee, main committee, society) for review and voting. This whole process may take more than a year. Everyone is free to join ASTM and take part in standard development. ISO standards are developed by independent technical experts who elaborate a draft version. Then a voting process begins and continues until a consensus is reached. ISO standard development takes about three years from the first draft to the published version. Thus, in case of a new type of implant, it may take years until a suitable standard test method is available. To overcome this problem, it may be appropriate to test implants according to custom testing procedures. These procedures should take into consideration the loads that are expected to act on the device in vivo. Loading a spinal implant in vivo is very complex and is rarely heard of. In an attempt to improve repeatability and increase reproducibility, standard test methods greatly simplify this complexity. This simplification is one of the main disadvantages of standard testing because the results do not directly reflect the situation in vivo. Therefore comparative data are almost always necessary, which means that the mechanical safety and effectiveness of a new medical device must be compared to a well-known and successfully used implant. In the United States, this process is called a 510(k) submission or Premarket Biomechanics of the Spine. https://doi.org/10.1016/B978-0-12-812851-0.00013-6 © 2018 Elsevier Ltd. All rights reserved.

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Table 1  Advantages and Disadvantages of Preclinical Implant Testing According to Standard Test Methods Advantages

Disadvantages

– Standard testing procedures are internationally known and widely accepted.

– No standards are available for new and innovative implants. Because drafting and approval of a new standard can take years, standard development always lags behind implant development. Innovative implants therefore often have to be tested according to custom, implant-specific testing procedures. Such testing can be of high quality but usually requires more efforts to convince the regulatory bodies. – Easy test conduction is mostly associated with strong simplifications of the real loading in vivo. Therefore interpretation of the test results is difficult. For that reason, commonly comparative data is required from a predicate device. – Some standards leave flexibility for the user, which again reduces the interlaboratory comparability.

– Standard testing can easily be conducted.

– The interlaboratory comparability is increased.

Notification. But such comparative data are also often required for CE Marking. This comparison must demonstrate that the device to be marketed is at least as safe and effective (i.e., substantially equivalent to) as the legally marketed device (predicate device). On the other hand, the biggest advantage of the standard testing methods is their international recognition (Table 1).

­ VERVIEW OF CURRENT AVAILABLE STANDARDS FOR TESTING SPINAL O DEVICES The standards for preclinical testing of implantable devices for spinal surgery are numerous. Table 2 shows the latest revisions of all the current available documents with a short description of the main peculiarities of each standard. ISO and ASTM are responsible for designing and issuing the standards, with slightly different approaches, and some general comments can be drawn from the analysis of the standards, as follows: – ISO standards tend to give a complete description of the testing condition that must be applied, including the geometry of the setup, the entity of the loads to be applied, and in the case of dynamic tests, the run-out cycles; however, performance levels for the tested devices are not always given (e.g., the maximum amount of debris that can be accepted during wear tests is not defined). – ASTM standards generally denote methods for running tests for comparison purposes; this is a main requirement of the FDA for the introduction of new devices on the US market. In this light, some details (e.g., regarding the definition of the load to be applied in fatigue testing) are missing, and iťs up to the manufacturer to justify the test protocol and to demonstrate the equivalence of the new device with the predicates. – In the latest revisions of the standards, some cross-references between ISO and ASTM standards designed for testing similar devices are reported (e.g., in ISO 12189), often to give evidence of

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the differences between the two approaches and to explain under which conditions one should use one standard rather than the other. – In the latest revisions of the ASTM standards, some reference tables containing results of static tests on predicate devices were introduced to enhance the importance of a statistical approach; however, these data can also be used to compare the performances of the new device with predicate device performances. – Useful description of the rationale underlying the choice of the prescribed setup and bibliographic references are always reported in both ASTM and ISO standards.

­STUDY CASES I­NTERPRETATION OF CURRENT STANDARDS WITH FEM METHODOLOGY (APPLICATION TO ASTM F 1717/ISO 12189) The assessment of the fatigue performances of a screw-rod system in the preclinical phase is mainly based on the use of two standards, ASTM F 1717 and ISO 12189. Both standards simulate a 2-FSUs application, but whereas the former uses a vertebrectomy model without any anterior support, the latter uses calibrated springs interposed between each couple of vertebra to simulate an intervertebral disc (Fig. 1). As shown in Table 2, the ASTM standards do not specify the load to be applied, and the maximum load to reach the run-out has to be determined by the user. This engineering-based approach generally leads to finding and applying a load so low that it is hardly comparable to clinical use (Villa et al., 2014a, b). On the other hand, the ISO standards impose a maximum load of 2 kN until the same run-out is reached. Comparison of the results from the two different procedures may present difficulties, perhaps some that are impossible to overcome. Differences in the applied load and the typology of the anterior support (absent or of different stiffness) will produce different internal states of stress in the implant that, in turn, will strongly affect their fatigue behavior. Moreover, as much as possible, the tesťs results should predict the device’s behavior after it is implanted in the patient. This is generally accomplished if the geometry of the setup, the loads that are applied, and the number of cycles are representative of the ones the device will be subjected to in  vivo. In this light, we wondered, on the one hand, whether the ASTM experimental configuration is correctly representative of the physiological anatomy of the thoracolumbar spine and, on the other hand, whether the internal loads arising in the rods using the two procedures are representative of those that arise when everyday life activities are carried out.

­Analysis of the Anatomical Parameters Influencing ASTM F1717 Results

In regard to the suitability of the ASTM standard to propose a methodology that accurately represents the anatomy of the spine, an investigation was conducted to (i) compare the set of geometrical parameters imposed by the standard with their physiological value in the thoracolumbar tract; and (ii) determine a worst-case scenario for the stress state in the instrumentation deriving from the superposition of all parameters (La Barbera et al., 2014). The values of the anatomical parameters of the thoracolumbar tract were extrapolated from both literature data and direct measurements on patients subjected to stereoradiographic X-ray imaging

Table 2  Current ASTM and ISO Standards for Preclinical Evaluation of Spinal Devices With Details of the Test Protocol Standard— Year of Release

Year of Reapproval

Title

Tested Implant

Mechanical Evaluation

Wear of total intervertebral spinal disc prostheses. Impingement-wear testing and corresponding environmental conditions for test of lumbar prostheses under adverse kinematic conditions. Wear of total intervertebral spinal disc prostheses. Nucleus replacements. Wear of total intervertebral spinal disc prostheses. Loading and displacement parameters for wear testing and corresponding environmental conditions for test. Fatigue test method for spinal implant assemblies using an anterior support. Standard test method for static, dynamic, and wear assessment of extra-discal single level spinal constructs.

IVD prosthesis

Wear

IVD prosthesis— nucleus replacements IVD prosthesis

Fatigue, wear

Posterior stabilization devices Extra-discal constructs (motion preserving, rigid fixation devices)

Fatigue

Standard guide for mechanical and functional characterization of nucleus devices. Standard practice for static and dynamic characterization of motion preserving lumbar total facet prostheses. Standard test methods for occipital-cervical and occipital-cervical-thoracic spinal implant constructs in a vertebrectomy model.

Nucleus devices

Static, fatigue

Total facets prostheses

Static, fatigue

Occipital-cervical, occipital-cervicalthoracic constructs

Static, fatigue

ISO 1819217, part 3

Draft

ISO 1819210, part 2 ISO 1819208, part 1

–

ISO 12189-08 ASTM F2624-12

2011

ASTM F2789-10 ASTM F2790-10 ASTM F2706-08

2015

ASTM F2694-07

2016

Standard practice for functional and wear evaluation of motion-preserving lumbar total facet prostheses

Components for total facets prostheses

Wear

ASTM F2423-05

2016

Standard guide for functional, kinematic, and wear assessment of total disc prostheses.

Total disc prosthesis

Wear, fatigue

ASTM F2346-05 ASTM F2267-03

2011

IVD devices

Static, fatigue

Intervertebral bony fusion devices

Static

ASTM F2193-02

2014

2014 2013

Standard test method for evaluating the static and fatigue properties of interconnection mechanisms and subassemblies used in spinal arthrodesis implants.

ASTM F1717-96

2016

Standard test methods for spinal implant constructs in a vertebrectomy model.

Components for spinal fixation (screws, hooks, plates, rods) Intervertebral body fusion devices Interconnection mechanisms and subassemblies for spinal arthrodesis Posterior fixation devices

Static, fatigue

ASTM F2077-00 ASTM F1798-97

Standard test methods for static and dynamic characterization of spinal artificial discs. Standard test method for measuring load induced subsidence of intervertebral body fusion device under static axial compression. Standard specifications and test methods for components used in the surgical fixation of the spinal skeletal system. Test methods for intervertebral body fusion devices.

a

2011

2016

2014 2017

2011

Wear

Static, fatigue,wear

Static, fatigue Static, fatigue

Static, fatigue

AC, axial compression; AR, axial rotation; AT, axial torsion; B, bending; CB, compression bending; CS, compression shear; FE, flexion-extension;

Simple Motions Mode (a)

Simple Loading Mode (a)

Combined Loading Mode (a)

Environmental Conditions

Run-Out Cycles (millions)

1

FE, LB, and AR (assigned)

AC (assigned)

F + LB

Yes

1

Cervical, lumbar Cervical, lumbar

1

FE, LB, and AR (assigned) FE, LB, and AR (assigned)

AC (assigned), CS AC (assigned), CS

FE + LB +  AR + AC + CS FE + LB +  AR + AC + CS

Yes

10

Yes

10

Cervical, thoracolumbar Lumbar

1, 2 or more 1

–

CB (assigned)

–

Yes

5

FE, LB, and AR

S, CB, TB

–

–

Kinematics, subsidence –

Lumbar

1

FE, LB, AR

AC, AT, S, B

–

Yes

10 or 5 (for motionpreserving or fixation devices) 10

Lumbar

1

F, E, LB, AR

S, CB,TB

–

Yes

10

–

Occipitalcervical, occipitalcervicalthoracic Lumbar

2

AR

CB, TB

–

Yes

5

1

AC (indicative range)

LB + AR + A, FE + LB + A

Yes

10

Kinematics

Cervical, lumbar

1

AC (indicative range)

–

Yes

10

–

Not specified

1

F, E, LB, AR (indicative ranges) FE, LB, AR (indicative ranges) –

AC, CS

–

–

Yes

Subsidence

Not specified

–

–

AC

–

–

–

–

Not specified

–

–

B, AT

–

–

2.5

–

Not specified

1

AR

AC, CS

–

–

5

Loosening

Not specified

–

–

AC, AT, S, B

–

–

2.5

–

Cervical, thoracolumbar

2

AR

CB, TB

–

Yes

5

Functional Evaluation

Spinal Region

Number of FSUs

Impingement

Cervical, lumbar

– –

– –

Kinematics

1

FSU, functional spinal unit; LB, lateral bending; PE, polyethilene; S, shear; TB, tension bending.

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FIG. 1 ASTM F1717 (top) and ISO 12189 (bottom) experimental setups: anatomical sketch (A), sketch taken from the standard (B), specimens during the tests (C).

(Fig. 2); in parallel, a parametric finite element model of the ASTM F1717 setup was built and discretized, assuming its linear elastic material properties (Fig. 3). The simulations were run by imposing a vertical load of 300 N and using a systematic approach so that each anatomical parameter investigated varied between its minimum and maximum values, while all the other parameters were kept fixed. Stresses on the screw head and on the rod in the transverse plane were calculated this way, and results were presented in terms of a von Mises stress increase normalized on those calculated for the reference configuration proposed by ASTM. Fig. 2 shows an example of the results obtained for a particular parameter (the Block Moment Arm, BMA).

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BMA

BMA (mm) 15 T1

20

25

30

35

40

229

BMA

45

ASTM F1717

T2 T3 T4 Follower load path

T5 T6 Spinal level

T7 T8 T9 T10 T11 T12 +7%

L1 L2 L3 L4 L5 S1 –60 –50 –40 –30 –20 –10 0 sVM increase on the screw (%)

10

FIG. 2 BMA as a function of the thoracolumbar level: comparison between the value proposed by ASTM F1717 and data measured on healthy subjects. BMA is the lever arm of the eccentric load in ASTM setup (right), interpreted as the lever arm of the follower load in a healthy subject’s spine (left). The lower horizontal axis reports the percentage of stress increase on the pedicle screw for each value of BMA compared to ASTM reference configuration.

A further step was taken to identify the worst-case scenario for each spinal level, which meant identifying the combination of the different parameters’ value that gave the highest increase of stresses in the instrumentation. Also, in this case, a comparison of the state of stress arising when ASTM proposals are followed is presented (Fig. 4). Results showed that two anatomical parameters had the strongest influence over the stress arising in the instrumentation: the previously mentioned BMA and the center of fixation to rotation (CoFR). The latter is the vertical distance between the screw insertion point in the PE block and the center of rotation of the cylindrical pin used to apply the load in the ASTM setup, whereas in the anatomical model, it is the distance between the screw insertion point and the instantaneous center of rotation of the FSU. The results demonstrate that the setup proposed by ASTM represents well the stress on the device when it is implanted at the L1 level in an average patient taken from a physiological population (Fig. 4). However,

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ISO 12189

ASTM F1717

k (N/mm) 0 Black

100 Green

147 Blue

375 Red

459 Yellow

k Grey

FIG. 3 FE model of ISO 12189 standard (top): calibrated springs with increasing axial stiffness (k) and different color codes (green, blue, red, and yellow) from left to right are used to mimic the support provided by the anterior spine, potentially describing a wide range of instability scenarios (bottom) from a full corpectomy to a rigid titanium cage. The ASTM F1717 model can be considered as a limit model without springs. Only one quarter of each setup was described, assuming its symmetry in the horizontal and sagittal planes.

FIG. 4 Results from FE analyses. The solid line indicates the average case (combination of the physiological mean values for each parameter); the dashed line indicates the worst case (combination of physiological worst values for each parameter).

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things drastically change when, at the same level, the upper boundary of the anatomical parameters of this population is considered. In such a patient, stress in the instrumentation can increase more than 15% on the screw, suggesting a possible revision of the setup that is able to take into account the actual worst-case scenario (La Barbera et al., 2014). A more recent experimental and numerical study investigated the effect of such a revision on a spinal fixator already cleared for commercial use. It was shown that the revised ASTM standard led to a significant 3.2-million-cycle drop in the fatigue performance compared to that of the same implant tested according to current ASTM recommendations (La Barbera and Villa, 2017).

­Analysis of the Anatomical and Mechanical Parameters Influencing ISO 12189 Results

A similar numerical analysis was performed on the ISO 12189 standard to determine the effect of mechanical and anatomical parameters on the loads occurring on the posterior stabilization device (La Barbera et al., 2016a; La Barbera and Villa, 2016). An FE model of the ISO standard setup was built, and a simple vertical load up to 2000 N was applied on the construct (Fig. 3). The stiffness of the calibrated springs could be varied among few standard axial stiffness values, whereas the ASTM model represented a limited case where stiffness is set to zero. A stepwise validation procedure was developed to ensure the reliability of the FE predictions for the experimental data (La Barbera et al., 2015; La Barbera and Villa, 2016). Simulations demonstrated the primary role played by the anterior support stiffness and its position on the sagittal plane (namely, the anterior support arm, or ASA) over other anatomical parameters. Spring stiffness alone contributed a stress increase up to 360%, whereas BMA reached 10% (La Barbera et al., 2016a). A worst-case combination of anatomical parameters could lead to a maximum stress increase of 22.4% at L1, compared to the current ISO standard.

­Analysis of the Suitability of Standards to Represent Daily Activities

In order to study the internal loads arising in the posterior implant during ISO and ASTM standard testing, an FE study was conducted (La Barbera et al., 2016b, 2017). Results were compared to the predictions obtained on an average L2-L4 spine segment that, through corpectomy, had been instrumented with a posterior fixator, a physiological anterior support, and a rigid anterior cage (Fig. 3). Moreover, data were compared to published in vivo measurements of real patients performing everyday activities. Numerical predictions were represented as a function of the vertical load and were compared to published in vivo measurements taken on patients bearing instrumented spinal fixators as they performed simple daily activities, as shown in Fig. 5 (Wilke et al., 2001; Rohlmann et al., 1994, 1997, 2000). Results of all the simulations are presented on simple graphs that make it easy to compare the internal actions in the rods calculated with the FE models to the ones measured in the patients. These graphs can be used in many ways. Here are some examples that refer to the calculation of the bending moment arising in the rods: (i) Case A in Fig. 6—To verify which daily activity produces stresses in the rods closer to the ones arising in the ISO in vitro experiments and when the suggested springs (red) and vertical load (200 N) are used. In this case, we found that the ISO setup reproduces quite well a 2-FSUs stabilization of a real patient, as well as of an average lumbar spine, undergoing flexion of the upper body.

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Standing

–3.9

–4.5

–3.6

–4.3

–4.3 –4.9

–4.2

After the implant Long term

–4.6

–3.4 –5.1

–2.0

–1.5

–3 –6 –12

–9

Moment (Nm)

0

3

In vivo measurements

Walking

Flexion

Extension

FIG. 5 Results from direct in vivo measurements performed by Wilke et al. (2001) and Rohlmann et al. (1994, 1997, 2000).

FIG. 6 Results of Case (A), case (B), case (C), and case (D). Please refer to the text for the complete explanation of the four cases.

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(ii) Case B in Fig. 6—To identify which vertical load has to be applied, again using the suggested red springs, for example, when standing is the activity to be simulated. In this case, the suggested value for the vertical load is about 1600 N. (iii) Case C in Fig. 6—To identify which typology of springs must be used when we want to simulate a walking patient. In this case, it was not possible to find a value for the vertical load to be applied to a construct that used red springs, and blue springs can be adopted for a 1200 N vertical load. (iv) Case D in Fig. 6—To identify which load to apply to an ASTM construct that gives the same stress state as the ISO one when standard proposals are used (red springs, 2000 N vertical load), or to associate the state of stress deriving from the application of a vertical load to an ASTM setup to a particular daily activity. This result is particularly useful for comparison purposes, because results of fatigue testing of predicate devices are often available only according to the ASTM standard. Such a testing method is not representative of a relevant clinical condition, because vertebrectomy is generally associated with the presence of an anterior support. Thus the graph allows matching these available results to real clinical data measured during the daily activities of representative patients.

I­MPLEMENTATION OF A NEW STANDARD PROCEDURE: TESTS FOR EVALUATION OF SCREW LOOSENING Mechanical failure of a posterior spinal fixation device happens from time to time. Cases of broken screws or rods occur, and in other cases, the polyaxial head of the pedicle screws or the set screws loosen. Therefore testing according to several standard procedures is required before a fixation device can enter the marketplace. Whereas testing according to ASTM F1717 includes the whole fixation system, ASTM F1798 focuses on subassemblies, such as one rod plus one screw to which the rod is attached. ASTM F2193 finally describes how to test single components of a fixation device. A single screw, for example, can be tested for its bending characteristics. None of these standards, however, take into account anchorage of the screw onto the vertebral bone. From a clinical point of view, this anchorage plays an important role because screw loosening is a common clinical complication requiring surgical intervention (Galbusera et al., 2015). To overcome this clinical complication, new and innovative pedicle screws have been developed. Some screws are expandable; others are equipped with innovative thread designs or are cannulated for injection of bone cement or other materials. The only standardized possibility to investigate the effectiveness of the anchorage between a screw and a bone is to conduct a pull-out test according to ASTM F543 (Fig. 7). According to this standard, the screw is drilled into a bone substitute, usually polyurethane foam, and then pulled out. In vivo, however, a screw not only experiences pull-out loads. Actually, a screw that is part of a posterior spinal fixation device also experiences repetitive bending and shear loading. This loading condition is not reflected in the standard pull-out test (Sandén, 2010; Costa et al., 2013; Kueny et al., 2014). Therefore, in the past, some researchers have published more physiological test setups (Zindrick et al., 1986; Law et al., 1993; Tan et al., 2004; Kueny et al., 2014). Some of these setups are called toggle-tests. In this type of test, the screw, which is implanted into a cadaveric vertebra, is loaded cyclically in cranio-caudal direction. This loading condition is already much closer to reality than a static pull-out test. However, standardization is difficult as long as testing is carried out by using cadaveric specimens. Cadaveric specimens are scarce, and the available ones are mostly from elderly

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FIG. 7 Sketch of the pull-out setup according to ASTM F543. The screw is drilled into a vertebral body substitute (polyurethane foam) and then pulled out by application of a continuously increasing load.

donors, and they usually do not represent the age of a typical patient. A standard test method, however, needs to be reproducible and conductible under the same conditions. For that reason, a mechanical test setup needed to be developed that does the following: – – – – – – –

Allows testing of various screw designs and anchorage systems. Mimics the in vivo loading situation at the screw-bone interface as closely as possible Is validated. Can be conducted easily. Can be conducted on any uniaxial materials testing machine. Is independent of the testing laboratory to the greatest possible extent. Can be conducted economically.

Based on these requirements, the new setup (Fig. 8) consists of a vertebral body substitute made of polyurethane foam (Open Cell Block 15 PCF, Sawbones, representing healthy cancellous bone). Additionally, a pedicle substitute made of carbon fiber-reinforced epoxy resin was designed. Its inner and outer geometry was adapted to those of the human lumbar pedicles (Zindrick et al., 1987). This pedicle substitute was needed to allow the pedicle to deform elastically as it presumably does in vivo. The vertebral body and pedicle substitutes were rigidly clamped to each other. The screw was drilled through the substitute pedicle into the substitute vertebra.

­STUDY CASES

Pedicle fixture Pedicle

235

Vertebral body fixture Cover plate

Vertebral body (PU foam)

FIG. 8 Sketch of the vertebral body substitute made of polyurethane foam, which is clamped to a pedicle substitute made of a carbon fiber-reinforced epoxy resin.

FIG. 9 New test setup, which incorporates shear forces, bending moments, and pull-out forces. The relationship between these three loading components is based on the literature.

Then a rod was fixed to the head of the screw. The rod was rigidly fixed to the bottom plate of the materials testing machine under an inclination of 27 degrees (Fig. 9). This inclination had been used in other standard testing procedures (ASTM F2077) and was selected to generate a physiological relationship between shear load, bending moment, and pull-out force (Rohlmann et al., 1994, 2000). The load was applied to the vertebral body substitute using hinge pins around which the vertebral body was allowed to rotate during testing. The validation of the vertebral body and pedicle substitutes was based on a comparative study. Six screws (S4, Aesculap) were tested using the new substitutes in the new test setup as previously described, and another six screws were tested using the new test setup up but with cadaveric vertebral bodies (mean age 58 ± 7  years, mean bone mineral density 75.4 ± 4.0 mg/ccm, 4 × male, 2 × female) instead of the vertebral body and pedicle substitutes. The load was increased stepwise every 50,000 loading cycles until significant loosening was observed. The first load step consisted of a 100 N inferior shear load, a 50 N pull-out load, and a 3.8 Nm flexion moment.

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FIG. 10 Flexion angle between screw and vertebra at each load level for each of the six test runs with the vertebral body and pedicle substitutes and with the cadaveric specimens.

The results showed that there was less variability between the displacement curves of six test runs in the group with the vertebral body and pedicle substitutes (Fig. 10). Also, the displacement in these six test runs was within the range of the displacement of the six test runs with the cadaveric specimens. However, a detailed analysis of the relative motions between the screw and the vertebra revealed that the screw never loosened at its entry point into the pedicle in the substitutes, whereas it sometimes loosened in the cadaveric specimens. In conclusion, the new vertebral body and pedicle substitutes allow more standardized and reproducible testing. In combination with the new setup, which allows application of shear, bending and pull-out loads at the same time, the substitutes may have the potential to become an easily conductible and reproducible testing procedure for investigation of the anchorage between screw and vertebra. In the future, further comparative tests with the various new screw designs will be needed to prove their general usability.

­CONCLUSIONS/FUTURE TRENDS Notified bodies request standards during experimental preclinical testing of a spinal device. Despite simplifications of the real in vivo conditions are necessary to improve comparability and repeatability of each standard method, specific aspects deserve a much deeper analysis. Standards should be continuously questioned, and an increasing awareness needs to be developed in order to have clinically relevant testing procedures that are really predictive of in vivo performances.

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