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93.Verdonschot N, Huiskes R. Creep properties of three low temperature-curing bone cements: a preclinical assessment. J Biomed Mater Res 2000;53B:498–504.

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98.Harris WH. Is it advantageous to strengthen the cementmatel interface and use a collar for cemented femoral components of total hip replacement. Clin Orthop Rel Res 1992;285:67–72.

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See also BIOMATERIALS, TESTING AND STRUCTURAL PROPERTIES; HIP JOINTS, ARTIFICIAL; ORTHOPEDICS, PROSTHESIS FIXATION FOR; RESIN-

BASED COMPOSITES.

BONE DENSITY MEASUREMENT

YIXIAN QIN

ERIK MITTRA

Stony Brook University

New York

INTRODUCTION

Chronic diseases, such as musculoskeletal complications, have a long-term debilitating effect that greatly impacts quality of life. Osteoporosis is a reduction in bone mass or density that leads to deteriorated and fragile bones and is the leading cause of bone fractures in postmenopausal women and in the elderly population for both men and women. About 13–18% of women aged 50 years and older, and 3–6% of men aged 50 years and older, have osteoporosis in the United States alone. These rates correspond to 4 6 million women and 1 2 million men who suffer from osteoporosis (1). One-third of women over 65 will have vertebral fractures and 90% of women aged 75 and older

have radiographic evidence of osteoporosis (2–4). Another 37–50% of women aged 50 years and older, and 28–47% of men of the same age group, have some degree of osteopenia. Thus, approximately a total of 24 million people suffer from osteoporosis in the United States alone, with an estimated annual direct cost of over $18 billion to national health programs. Hence, early diagnosis that can predict fracture risk and result in prompt treatment is extremely important. Early identification of fracture risk, most commonly caused by osteoporosis-induced bone fragility, is also important in implementing appropriate treatment and preventive strategies. Indeed, the ability to accurately assess bone fracture risk noninvasively is essential for improving the diagnostic as well as therapeutic goals (i.e., assessing temporal changes in bone during therapy) for bone loss from such varied etiologies as osteoporosis, microgravity, bed rest, or stress-shielding around an implant.

Assessment of bone mineral density (BMD) has become as essential element in the evaluation of patients at risk for osteopenia and osteoporosis (2,5–8). Bone density was initially estimated from the conventional X ray by comparing the image density of the skeleton to the surrounding soft tissues. Although demineralized bone has an image density closer to soft tissues, dense mineralized skeletal tissues appears relatively white on an X-ray image. Hence, the mineral density of bone can be estimated by the degree of gray color of the X-ray image in the bone region. However, because of its resolution and variations generated in the X-ray image, it has been suggested that bone mineral losses of at least 30% are required before they may be visually measured using a conventional X ray (9,10). Growing awareness of the impact of osteoporosis on the elderly population and the consequent costs of health care, together with the development of new treatments to prevent fractures, have led to a rapid increase in the demand for bone densitometry measurements. Many image modalities and techniques have been developed to improve the quality and the accuracy of the measurement for bone mineral and dense assessment. Two major densitometry techniques are commonly used in assessing bone density, that is, radiography-based densitometry and ultrasoundbased assessment.

RADIOGRAPHY-BASED DENSITOMETRY

To improve the sensitivities of X-ray images to bone density changes and assessment, several technologies have been developed. Bone densitometry is a term that is defined as a method for imaging density of bone. However, the ‘‘true’’ density is not applicable in the current radiography-based techniques. In the field of densitometry, the term ‘‘bone mineral density’’, referred to as BMD, is related to the mass of bone in the tissue level, which includes both bone and marrow components as well as surrounding soft tissues. Furthermore, most densitometric techniques are projectional for the image formation that provides a twodimensional image of the three-dimensional (3D) bone volume being measured. Therefore, the BMD defined from the projectional techniques is the mass of bone tissue mass (including marrow and/or soft surroundings) per unit area

in the image, not per unit volume of the tissue. Hence, what is actually measured is the apparent bone mineral density, which is defined by the bone mineral content contained in the area scanned, or expressed as gram per squared centimeter in unit. To detect osteoporosis accurately, several methods are developed for the noninvasive measurement of the skeleton for the diagnosis of osteopenia, osteoporosis, and/or the evaluation of an increased risk of fracture (11). These methods include single-energy X-ray absorptiometry (SXA), dual energy X-ray absorptiometry (DXA or DEXA), quantitative computed tomography (QCT), peripheral quantitative computed tomography (pQCT), radiographic absorptiometry (RA), dual photon absorptiometry (DPA), and single photon absorptiometry (SPA). There are two types of BMD measurements, peripheral BMD and central BMD. The peripheral BMD instruments are usually smaller, less expensive, and more portable than the central BMD. Central BMD is capable of measuring multiple skeletal sites, that is, the spine, the hip, and the forearm. Table 1 lists these methods currently available for the noninvasive measurement of the skeleton for the diagnosis of osteoporosis. These techniques differ substantially in physical principles, in the particular physical body sites (e.g., spine, hip, or total body), in the clinical discrimination and interpretation, and in availability of the facility and cost.

Single-Energy Densitometry

This instrumentation passes a beam of radiation through the limb of the body (e.g., forearm) and determines the difference between the incoming (or incident) radiation and the outgoing (or transmitted) radiation, referring to the attenuation. The higher the bone mineral content, the greater the attenuation. Mineral content can be calculated by the attenuation of the radiation. BMD can then be calculated by dividing the mineral content by the detected bone area. The relation between incoming and outgoing X- ray energy can be expressed as

where I0 ¼ incoming radiation intensity, I ¼ transmitted radiation intensity, l ¼ mass attenuation coefficient, and d ¼ area density of the attenuating materials (g cm 2). The mass attenuation coefficient is a physical property that describes how much a given material attenuates and X-ray energy. If the attenuation coefficient can be

experimentally determined, the equation becomes explicit, and the area density can be determined by

where k is an experimentally determined constant for the attenuation coefficient.This technology is relatively simple and easy to understand. However, biological tissues and body are composed by multiple materials (e.g., bone, muscle, and other soft tissues). The accuracy of the technology is limited.

Single-Photon Absorptiometry

Bone density can be measured by passing a monochromatic or single-energy photon beam through bone and soft tissue. This procedure is referred as SPA. The amount of mineral content can be quantified by the attenuation of the beam intensity. After the photon beam attenuation is calculated, the value of the attenuation can be compared with a calibration parameter derived from a standard mineral content (e.g., using ashed bone of known weight). This procedure can finally determine the BMD with measured attenuation. Iodine-125 at 27 keV, or americium-241 at 59.5 keV, was initially used for generating of the SPA beam. SPA is rarely used in clinical practice today. SPA determined bone mineral content is calculated through uniform thickness of the soft tissue in the path of the beam. The targeted scan site (e.g., limb or forearm) had to be submerged in the water or a tissue-equivalent material, which limited the practical applications of the SPA. The advantages of this technique include a low dose of radiation, portable, and use for particular body sites with relatively precisely measurement. Although the SPA is an approximate method, the limitations of SPA include limited accuracy of the measurement, radiation, and used only on the particular peripheral sites, like forearm and heel.

Dual-Photon Absorptiometry

To overcome the limitations of single-energy or photon densitometry, if a dual radiation source was used, the influence of soft tissues could be eliminated. The basic principle involved in DPA for bone density measurement was similar to SPA. The degree of attenuation of the photon energy beam between incoming and outgoing energy through bone and soft tissue is quantified. As with SPA, the beam source was originally used, but with an isotope

552 BONE DENSITY MEASUREMENT

used, which emitted photon energy at two distinct photoelectric peaks. When the beam was passed through a region of the body with both hard and soft tissues, attenuation of the photon beam appeared to both photon energy peaks. The contributions of soft tissue to beam attenuation can be determined by the quantifications of the relative relations between two attenuations (15). Because of its capability to distinct bone from soft tissue, DPA has been used to quantify bone density in deep tissue and large skeletal areas where bone is surrounded by large volume of soft masses (e.g., spine and hip) (16). DPA was considered a major advance from SPA due to its ability to quantify BMD and mineral content in such deep areas like spine and hip, as well as its capability of quantifications of effects from soft tissues. However, DPA has many notable limitations. First, the maintenance of the beam source was expensive, which had to be replaced yearly. Second, the radioactive source decay increased as much as 0.6% per month, which added difficulties for the calibration. These factors may result in the precision of 2–4% for DPA measurements in the region of interests. This precision (e.g., 2%) would limit its clinical application, in which a great change (e.g., 5–6%) from the baseline value had to be observed before one could reach the 95% confidence level for the change of bone density. Nevertheless, the concept of dual-photon densitometry has impacted the development of new technologies such as DXA.

Dual-Energy X-ray Absorptiometry

Perhaps the most popular bone densitometry used in clinical practice is the DXA or DEXA. The basic principles of DXA are the same as DPA. To overcome the major limitation of DPA, it did not take long for manufacturers who originally had the DPA product to replace the decaying isotope beam source with a highly stable dual-energy X-ray tube. There are several advantages of using X-ray sources over radioactive isotopes (i.e., no beam decay concerned in the X-ray tube and no calibration required for correction of the drifting because of the source decay in the DPA). The fundamental basis for DXA is the measurement of the transmission through the body of X rays of two different photon energies. The radiation source is collimated to a pencil beam and aimed at a radiation detector placed directly opposite the objective to be measured (Fig. 1). The patients are positioned on a table in the path of the X-ray beams. Due to the dependence of the attenuation coefficient on atomic number and photon energy, assessment of the transmission factors and attenuations at two energies enables the 2D apparent density, that is, bone density per unit projected area, of two different types of tissues to be inferred (17–19). The X-ray source and detector pair is scanned back and forth across the region of interests in the body, generating annotation images, which the BMD is calculated as the ratio of the bone content to the measured area. Radiation dose to the patients is very low on the order of 1–10 mSv. The DXA system can measure the BMD of the spine, proximal femur, forearm, and the total body. Recent technology uses a fan beam geometry in the DXA scanners that can increase the speed and reduce the acquisition time (GE Medical System Inc.). The image

Detector

Pencil beam

Figure 1. Scan path pattern for DXA densitometer using pencilbeam format.

quality of recent DXA has improved significantly via computational capability for better visualization. In addition to the body DXA scanners, recent technology has also adapted for development of lower cost, small, and particular site densitometries (Fig. 2). These systems are available to the clinic for specific body regions (e.g., spine, hip, leg, arm, and hand). The DXA systems are available for the diagnostic clinical use by many major manufacturers (e.g., GE Medical Systems of Madison, Norland, and Hologic Inc. of Bedford).

The basic working principle of DXA and its ability to reduce the effects of soft tissue is to use two X-ray sources and mathematically solve the bone thickness and softtissue thickness (15). By using two X-ray energies, two equations can be derived by scanning the measurement site twice with low (L) and high (H) energies once at each.

where I0 ¼ incoming radiation intensity, I ¼ transmitted radiation intensity, l ¼ mass attenuation coefficient, d ¼ area density of the attenuating materials (g cm 2), and b and s refer to the bone and soft tissue. Two scans are usually performed simultaneously with either two energies or rapid switching between two energies. When the

Figure 2. DEXA machine for a whole-body scan (QDR4500 fanbeam scanner, Hologic Inc., Bedford, MA) (left). DEXA bone densitometry is widely used in.

attenuation coefficients for bone and soft tissue are known for both low and high energies, the apparent or area bone mineral density can be calculated as

The attenuation coefficient for bone is relative constant but varied to persons. The soft-tissue attenuation coefficient, however, is contributed by fat and other soft tissues and varied greatly in the body. This is the source for the errors generated in the measurement. The manufacturers usually provide phantoms for calibration for the system.

Traditionally, the focus of clinical bone evaluation has been apparent or area BMD as measured by DXA or DEXA (20–23). DEXA provides an effective way to measure BMD in a specific region of interest and is the most widely used diagnostic modality for assessing osteoporosis and osteopenia (8,24–26). However, in particular, DEXA suffers from several shortcomings. Although density (quantity) does positively correlate with strength (27,28) and fracture risk (29–31), anywhere from 10–90% of the variability in bone strength remains unexplained the (32). Additionally, as discussed below, the stereology of trabecular bone is one of its distinguishing features (especially with respect to its mechanical behavior), but because DEXA provides only a 2D image of apparent density, it is inherently limited in this regard. DEXA also suffers from an inability to differentiate trabecular from cortical bone. Although it is true that cortical bone also deteriorates with age (33,34), the effects of bone loss are more prevalent in trabecular bone due to its much higher surface area, and the greater net amount of bone mineral content in cortical bone can conceal small changes in the trabecular bone when measuring only BMD. Nevertheless, as one of the key factors that contribute to the bone’s quality evaluation, BMD measured by DEXA is a most popular modality used in assessing the status of bone and the risk of fracture.

Quantitative Computed Tomography (QCT)

measurement with DEXA (35–43). Because of its high resolution, QCT can provide the measurement in the trabecular region (e.g., femoral neck and vertebral bodies) (39,40,44,45). Compared with DXA the advantage of QCT is the image-based cross-sectional anatomy, which allows for a selection of the region of interests (ROI) and a better assessment of geometrical properties (Fig. 3). Most CT systems provide a software package to automate the placement of the ROI within a particular body volume, e.g., vertebral bodies. QCT scans are generally performed using a single kilovolt setting (single-energy QCT). It is possible to use a dual-energy QCT, which can provide further improvement of the resolution, but at the price of poorer precision and higher radiation dose. New 3D volumetric techniques acquire datasets with which analysis of bone macroarchitecture may be further optimized. Due to its capability of high resolution, geometric and structural parameters determined in QCT may contribute to determine bone strength when integrated with other technology (i.e., finite element analysis). The advantage of spinal QCT is the high responsiveness of the vertebral trabecular bone to aging and disease, whereas the principal disadvantage is the cost of the equipment and the dosage received for the scanning (higher than DEXA).

Peripheral QCT (pQCT)

pQCT systems are available for measuring the forearm. The advantages of these devices are the capability of separating the trabecular and cortical bone of the ultradistal radius and of reporting volumetric density. Several clinical used pQCT devices are available (e.g., the Stratec XCT 2000, that are suitable for use in a physician’s office or in primary care).

QUANTITATIVE ULTRASONOMETRY

Quantitative ultrasound (QUS) for measuring the peripheral skeleton has raised considerable interest in recent years. New methods have emerged with the potential to estimate trabecular bone modulus more directly. QUS provides an intriguing method for characterizing the

material properties of bone in a manner that is noninvasive, nonionizing, nondestructive, and relatively accurate. The primary advantage of QUS is that it is capable of measuring not only bone quantity (e.g., BMD), but also bone quality (i.e., estimation of the mechanical property) of bone. Over the past 15 years, several research approaches have been developed to quantitate bone mass and structural stiffness using QUS (46–48). Preliminary results for predicting osteoporosis using QUS are promising, and it has great potential for widespread applications (including screening for prevention). As such, many QUS machines have been developed, and there are currently many different devices on the market. Most available systems measure the calcaneus using plane waves that use either water or gel coupling [e.g., Sahara (Hologic Inc., MA), QUS-2 (Metra Biosystems Inc., CA), Paris (Norland Inc., WI), and UBA 575 (Walker Sonix Inc., USA)] (Table 2). Recently, an image-based bone densitometry device for calcaneous ultrasound measurement is also made available using an array of plane ultrasound wave (GE-Lunar, Inc., USA). Using several available clinical devices, studies in vivo have shown the ability of QUS to discriminate patients with osteoporotic fractures from age-matched controls (49– 51). It has been demonstrated that QUS predicts risk of future fracture generally as well as DEXA (51–54). However, there are several noted limitations, including the tissue boundary interaction, the nonlinear function of density associated with bone ultrasonic attenuation, the single index covering a broad range of tissues (including the cortical and trabecular regions), and the interpolation of the results. Recently, a focused ultrasound sonometer device was developed to obtain the likelihood of a broadband ultrasound attenuation (BUA) image in the human calcaneus region (center frequency 0.5 MHz, focus 50 mm) (55,56) (UBIS 5000, Diagnostic Medical Systems; and DTU-one, Osteometer MediTech). These devices provide ultrasound images in the calcaneus region, in which the parameter compares with DEXA data. Perhaps the major drawbacks of these ultrasound osteometers are low resolution and lack of physical interrelation with meaningful bone strength. Although only showing the correlation between BUA data and BMD, these devices mostly provide qualitative information for assessment of osteoporosis, not the true prediction for bone structural and strength properties. Therefore, QUS remains at a stage as a screening tool (Fig. 4), because of the nonuniformity of the porous structure in the bone tissue and its associated effects in resolution (14). Research attention is focused on developing systems to provide true images reflecting the bone’s struc-

tural and strength properties at multiple skeletal sites, i.e., in the hip, which can provide a true diagnostic tool (instead of just for screening) that surpasses the radiation based DEXA machines.

If QUS bone densitometry can be developed to provide a ‘‘true’’ bone quality parameter-based diagnostic tool (i.e., directly related to the bone’s structural and strength properties) and to target multiple and critical skeletal sites (e.g., hip and distal femur), QUS would have a greater impact on the diagnosis of bone diseases (e.g., osteoporosis) than current available bone densitometry. Research efforts are made in this regard (55–59). As an example, a new QUS modality, called the scanning confocal acoustic diagnostic system, has been developed (57–60), which is intended to provide true images reflecting the bone’s structural and strength properties at a particular skeletal site at a peripheral limb and potentially at deep tissue like great trachanter. The technology may further provide both density and strength assessment in the region of interests for the risk of fracture (57–60).

Fundamental QUS Parameters in Bone Measurement

In an effort to use QUS for predicting bone quality, a variety of approaches have been explored with many studies published in the past decade, that have examined the utility of QUS and its potential application as a diagnostic tool for osteoporosis. The physical mechanisms of ultrasound applied to bone may include several fundamental approaches, [i.e., speed of sound (SOS) or ultrasonic wave propagating velocity (UV), sound energy attenuation

Figure 4. A QUS bone densitometry test in a heel region. Reproduced courtesy of GE-Lunar Inc.

(ATT), BUA, and critical angle ultrasound parameters] that closely relate to acoustic transmission in a porous structure. Most commonly, parameters for QUS measurement are BUA and SOS, which can be used to identify those persons at risk of osteoporotic fracture as reliably as BMD (52–54,61,62). It has been shown that both BUA and SOS are decreased in persons with risk factors for osteoporosis, that is, primary hyperparathyroidism (63–66), kidney disease (67), and glucocorticoid use (68,69). The proportion of women classified into each diagnostic category was similar for BMD and QUS. Using the World Health Organization (WHO) criteria to classify osteoporosis for BMD measurement using DEXA and QUS testing, approximately one third of postmenopausal women aged 50þ years with clinical risk factors were diagnosed as osteoporotic compared with only 12% of women without clinical risk factors. This suggests that the measurement of QUS with calcaneal BUA and SOS is to some extent the same as the BMD Z- score measurement.

Background of BUA in Trabecular Bone Measurement

BUA and SOS are currently two commonly used methods for QUS measurements, which make it potentially possible to predict bone density and strength. As an ultrasound wave propagates through a medium, BUA measures the acoustic energy that is lost in bone (unit: dB/MHz). The slope at which attenuation increases with frequency is generally between 0.2 and 0.6 MHz, and it characterizes BUA. The slopes of the frequency spectrum may reflect the density and structure of bone. Although relatively little is known about the fundamental interactions that determine ultrasound attenuation in bone, the potential sources contributing to the attenuation include absorption, scattering, diffraction, and refraction (70–73). Although absorption predominates in cortical bone attenuation, the mechanism of BUA in cancellous bone is believed to be scattering (14,74–76). The importance of scattering has been alluded to in the literature. Scattering is also suggested to contribute to the nonlinear variation in BUA with density observed in cancellous bone and a porous medium (77–79).

Background of SOS or UV for Bone Measurement

The strength of trabecular bone is an important parameter for bone quality. In vitro studies have correlated the ultrasound velocity with stiffness in trabecular bone samples (80–82). This indicates that ultrasound has the potential to be advantageous over the X-ray based absorptiometry in assessing the quality of bone in addition to the quantity of bone. The mechanism of SOS in predicting bone strength is believed to be due to the fact that the velocity of an ultrasound wave depends on the material properties of the medium through which it is propagating, but it also depends on the mode of propagation. By determining the wave velocity through a bone, the elastic modulus of bone specimens can be evaluated, or at least be approximated (80,83). When ultrasound travels through a porous material, e.g., trabecular bone, it carries information concerning material properties, such as density, elasticity, and architecture. A relationship exists between the ultrasound velo-

city (unit: m/s) and the material elasticity E and density

r (14,80)

p

The velocity with which ultrasound passes through normal bone is fast and varies depending on whether the bone is cortical or trabecular. Speeds of 2800–3000 m s 1 are typical in cortical bone, whereas speeds of 1550– 2300 m s 1 are typical in trabecular bone.

It is demonstrated that trabecular bone strength is highly correlated with elastic stiffness (84). With the introduction of QUS, several new diagnostic parameters and experimental results, both in vitro and in vivo, have shown potential for evaluating not only bone quantity (i.e., BMD), but also bone quality (i.e., structure and strength). Two principal variables, BUA and UV, have been confirmed to identify those persons at risk of osteoporotic fracture as reliably as BMD from DEXA. However, SOS and BUA are related to bone density and strength as well as to trabecular orientation, the proportion of trabecular bone and cortical shell, the composition of organic and inorganic components, and the conductivity of the cancellous structure. Thus, QUS of trabecular bone depends on a variety of factors that contribute to the measured ultrasound parameters.

Other Bone Status Measurement Methods and Motivation to Assess Bone Quality

Beyond bone quantity, the quality (the integrity of its structure and strength) has become an equally or even more important measure to understand the bone structure and mechanical integrity. Most osteoporotic fractures occur in cancellous bone. Therefore, noninvasive assessment of trabecular bone strength and stiffness is extremely important in predicting the quality of the bone. The strength of the trabecular bone mostly depends on the mechanical properties of the bone at the local and bulk tissue level, and on its spatial distribution (i.e., the microarchitecture). A better understanding of the factors that influence bone strength is a key to developing improved diagnostic techniques and more effective treatments. To overcome the current hurdles, to improve the ‘‘quality’’ of the noninvasive diagnostic instrumentations, and to apply the technology for future clinical application, new clinical modality may concentrate in several main areas: (1) increasing the resolution, sensitivity, and accuracy in diagnosing osteoporosis through unique methods for improvement of signal/noise ratio; (2) directly measuring bone’s strength as one of the primary parameters for the risk of fracture; (3) generating real-time compatible imaging to identify local region of interest; (4) validating structural and strength properties with new modalities; and (5) predicting local trabecular and bulk stiffness and microstructure of bone, and generating a physical relationship between measurement and bone quality. In an attempt to achieve these goals, recent advances of emerging technologies are developed primarily for animal studies at this stage. These include high resolution pQCT, micro-MR-derived measures of structure, micro-CT-based BMD, and combined assessment of strength using geometry, density, and computational simulation. These methods

556 BONE DENSITY MEASUREMENT

will further lead to a better understanding of the progressive deterioration of bone in aging populations, and ultimately they may provide early prediction of fracture risk and associated musculo-skeletal complications such as osteoporosis.

ACKNOWLEDGMENT

This work has been kindly supported by the National Space Biomedical Research Institute (TD00207 and TD00405 to Y. Qin) through NASA Cooperative Agreement NCC 9-58.

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See also BONE AND TEETH, PROPERTIES OF; COMPUTED TOMOGRAPHY.

are deeply involved in the process of fracture healing, and it has been suggested (2) that delayed healing may be the result of cessation of the periosteal response before bridging has occurred, while nonunion may be indicative of a breakdown of both the periosteal and the endosteal repair mechanisms. A more general term for nonunion is pseudarthrosis, or false joint. Worth noting is that this problem is also infrequently found at birth (congenital pseudarthrosis). Electric and electromagnetic treatment is prescribed for both types of psuedarthroses, those that are the result of ununited fractures, and those that are found at birth. Further, because spinal fusion following back surgery can be problematic, electromagnetic treatment is also being used as an adjunctive procedure to promote spine fusion (3).

THE ELECTRIC CHARACTER OF BONE

Bone has a number of remarkable physical properties, particularly its electric character. Its electrical properties and the intimate relation of these properties to the growth process in bone were brought to light in a series of experimental discoveries, beginning in the 1950s. These revealed

BONE UNUNITED FRACTURE AND SPINAL FUSION, ELECTRICAL TREATMENT OF

AR LIBOFF

Oakland University

Rochester, Michigan

DEFINING THE UNUNITED FRACTURE

Within hours following a fracture in bone and the rapidly resulting hematoma, an endogenous repair process is initiated, characterized by increased cell division in the periosteum and endosteal stem-cell differentiation leading to organization of the hematoma into fibrocartilaginous callus. The latter represents the source of osteogenic potential from which ossification and subsequent bone remodeling occurs. In the ideal case, with proper management, those suffering bone fractures will normally find themselves fully recovered within a few months, with this time varying according to the specific bone involved, the type of fracture, and the age of the patient.

However, a small percentage of fractures fall outside the norm and do not heal as readily. There are upward of 5 million fractures occurring each year in the United States

(1). Approximately 5–10% of these remain ununited after a few months. One can identify two types of ununited fractures, those undergoing delayed fracture healing, as evidenced by a lack of full healing in 3–6 months, and nonunions, where there is a lack of healing 6–12 months after the fracture has occurred. Marsh (2) suggests that the best measure of fracture healing in humans may be recovery of bending stiffness (i.e., the torque measured in Newtonmeters that will bend bone by 18). He defines delayed union as failure to reach a stiffness of of 7 N m deg 1 at 20 weeks following fracture. Both the periosteum and the endosteum

1. A piezoelectric effect in bone.

2.A striking bioelectric signature specifically associated with developing bone.

3.A characteristic signature in adult unstressed bone.

4.A characteristic bioelectric signature following bone fracture.

Piezoelectricity is the rather unique property in which mechanical force is transformed into electric polarization (Fig. 1). Bone was shown to be piezoelectric by Yasuda in the early 1950s, but the work leading to this conclusion was not made generally available until 1957(4). Fukada and Yasuda (5) later found that this property could be traced to the intrinsic collagen component in bone. Since that time, a number of observers (6–8) suggested that this mechanical stress–electric polarization property should more properly be referred to as a stress-generated potential (SGP), reflecting the fact that what actually happens may not be the result of the special sort of crystal or textural structure that underlies the piezoelectric effect, but might instead result from the well-known electrokinetic effect of streaming potential. Streaming potentials, similar to piezoelectric signals, are characterized by the transformation of mechanical stress into a potential difference. However, streaming potentials do not occur because of any intrinsic crystal structure, but rather because fluid displacement through porous materials or tubes results in electric charge separation. It is generally agreed that dry bone indeed exhibits piezoelectricity, but opinions vary on whether this effect actually plays a role when bone is in its usual (i.e., wet) physiological environment. Part of the difficulty in resolving this issue is that the piezoelectric effect is not easily measured in wet bone. Whatever the pros and cons concerning studies on wet bone, it is difficult to put aside the seminal experiment by McElhaney (9). More than 600 silver epoxy electrodes were attached to cover the surface of a dried intact human femur from autopsy, and a vertical

Figure 3. The BEP Profile. Bioelectric potential profile of a rabbit tibia (14). Voltage differences are obtained relative to the proximal end (on the left) by means of salt bridge electrodes.

negative with respect to the midshaft and distal portions of the bone. Because the BEP is unaffected by local nerve denervation or reduced blood flow, but slowly disappears following animal death, it is believed (17) that the origin of this voltage stems from functioning bone cells, acting in concert.

Other than this likely connection to bone cells, it is difficult to pin down a reasonable physical explanation for the ubiquitous potential profile associated with long bone. Electric polarization is readily observed in specimens of mature bone when they are even slightly heated. The origin of this effect is still unclear, but it may reflect a pyroelectric response having a textural origin (18), or perhaps, as Mascarenhas has suggested (19), bone is inherently an electret, a type of material, like many bioploymers, with the interesting property of being capable of storing electric charge. Electrets are the electrical equivalent of magnets, and some observers have suggested that bone exhibits ferroelectric properties. The characteristic property seen in electrets is a slow release of charge when heated. For example, long-term currents on the order of 100 fA can be observed (20) for bone specimens heated to 40 8C. Regardless of the cause, it is most likely the case, as stated by Brighton (21), that: . . .in living, non-stressed bone, areas of active growth...[are] electronegative when compared with less active areas.

There is one more impressive electric property associated with living bone, again reflecting this question of the role of negative potentials. Only a few hours following bone fracture, the bone becomes more negative relative to the prefracture BEP (22) (Fig. 4). There is some dispute as to whether this effect is limited to the fracture site or is distributed more widely along the length of the bone (14,23). This uncertainty is in all likelihood due to the fact that there are obvious measurement problems in obtaining a BEP profile for a fractured bone. As an injury current one might expect a more specific and localized expression. However, it is possible that the entire periosteum may be affected in a bone fracture at any point along its length. Worth noting are the experiments by Becker and Murray (24) on fracture healing in amphibian systems indicating a

Figure 4. Effect of fracture of an 8 cm rabbit tibia on BEP. Measured potentials are shown before and after the fracture, which is at the 4 cm point.

discrete electrical negativity at the fracture site, which led him to characterize the innate ability of bone to heal itself in higher animals as a form of regenerative healing.

Viewed in the context of its other electrical properties, the change in voltage profile associated with bone fracture has to be regarded as consistent with the overarching concept that bone makes extensive use of electricity in all of its growth, repair, loadbearing, and homeostatic processes. Because of this, it is hardly surprising that exogenous electric currents have been widely applied in attempts to grow and/or repair bone.

The FDA-approved devices for electric repair of ununited fractures fall either into invasive or noninvasive categories. The invasive devices make use of implanted direct current (dc) and (ac) electric signal sources, both pulsed and continuously sinusoidal. The noninvasive types are either purely electric (capacitive coupling or CC), or electromagnetic, using pulsed magnetic fields (PMF or PEMF) or ion cyclotron resonance (ICR) tuned magnetic field combinations.

DIRECT CURRENT OSTEOGENESIS

The surgeon who first observed that bone is piezoelectric, Iwao Yasuda, was also the first to demonstrate (25) that electric fields applied to long bone in vivo are capable of producing callus. He wrapped a few turns of wire around rabbit femur, and, maintaining this point at a negative potential, passed a small (1 mA) current to an anode located away from the bone. It was consistently observed that after 3 weeks this current resulted in spicules of osseus callus (called electric callus by Yasuda) (Fig. 5). Surprisingly, these spicules were not directed along the bone itself, but instead along the direction of the current, in some cases actually pointing away from the bone. To the orthopedic surgeon, one of the most positive signs during the course of fracture repair is the appearance of callus. Thus the observation by Yasuda cannot be overemphasized.

Battery Resistance

Figure 5. Formation of callus in response to 1 mA dc current. After original sketch by Yasuda (26), showing wires wrapped around the bone.

Although the reasons why electricity is capable of forming callus are still not clear, the implication of Yasuda’s work was that electrical stimulation might be of assistance in bringing ununited fractures to closure.

Most of the follow-up experiments to Yasuda’s discovery concentrated on determining the effects of electrical signals on normal and fractured bone. A commonly used animal experimental design was to apply an electrical signal to one femur while using the contralateral femur in the same animal as a control. Intrinsic to this approach was the use of dummy electrodes, carrying no current, but serving to affect the contralateral limb by its mere presence in whatever way the electrodes were affecting the activated side. It was in this manner that Bassett et al. (27) used implanted battery packs to deliver microampere-level currents to platinum–iridium wire electrodes extending into the medullary cavities of femora in dogs. The results clearly indicated that more bone was formed in the intermedullary space in the vicinity of the cathodes than near the anodes. Follow-up experiments (28) reinforced the finding that bone growth appeared to be effective at the cathode, but also that bone necrosis occurred at anodes for currents in excess of 20 mA. In this work Friedenberg et al. (28) found that bone growth was most pronounced for currents between 5 and 20 mA. There is some question concerning this optimal current in that the required levels may be dependent on the mode of application. In rabbit femur, circular defects 2.8 mm in diameter were repaired within 3 weeks when subjected to currents ranging between 2.5 and 3 mA applied by two electrodes on either side of the defect (29). Not only was the current lower than that suggested by Friedenberg et al. (28), but there was no particular advantage to either polarity. Similarly, Hambury et al. (30) studying 85Sr uptake in rabbit femur observed osteogenesis at 3 mA, again with no difference due to polarity. In another attempt (31) to establish the optimal current for repairing bone defects in dog, it was reported that 0.2 mA was more effective than either 2.0 or 20 mA.

Further complicating the issue of what level of current is required to initiate osteogenesis were a number of earlier reports in which callus was formed using currents that were orders of magnitude smaller than microampere

levels. Fukada and Yasuda (4) wrapped a charged Teflon electret around bone to initiate callus, work that was later successfully repeated in Japan (32,33). Three different types of current application were employed in the latter experiment: that emitted by an electret, that obtained from the piezoelectric poly-g-methyl-L-glutamate (PMLG) film, and a battery delivering 8–10 mA. The two current levels for the electret and the film, respectively, were 1 and 10 pA, levels smaller by huge factors of 10 7 and 10 6 from the ‘‘optimal’’ value of 10 mA. Marino and Becker (34) raised the issue as to whether this enormous difference in currents, both seemingly effective, means that more than one mechanism is involved, with the microampere (mA) results indicative of a nonspecific osteogenic stimulus while the picoamp (pA) currents more closely mimicing the endogenous piezoelectric response.

ELECTROMAGNETIC OSTEOGENESIS

Among his other important discoveries, Michael Faraday was the first to show that voltage is induced in a conductor when a nearby magnetic field is changing rapidly. This phenomenon, often referred to as Faraday’s law, can be mathematically expressed by the following expression:

where dB/dt is the time rate of change of the magnetic field B through a region of area A, and V is the voltage induced by dB/dt along the path that is circumferential to A by this rate of change. If B is varying at some frequency f, the product fB is a good measure of the relative effectiveness of dB/dt. When the region in question is electrically conducting, as in living tissue, one can use Ohm’s law to rewrite the above expression in terms of the current I instead of V. Thus if R is the resistance of the circumferential path around A, Eq. 1 is changed to read

In this way, one can induce a current in the vicinity of a bone defect by employing a nearby magnetic field that is changing rapidly—the faster the rate of change, the greater the current. One achieves a faster change (i.e., a larger dB/dt) by merely increasing the frequency at which B is changing. Further, it is important to realize that the current so induced is no different from currents that are produced by purely electrical means (Fig. 6). Most important, since the source of the magnetic field can be deployed externally, Faraday’s law enables the clinician to generate the required therapeutic currents in a completely noninvasive manner.

In 1974, following 5 years of intensive effort. Bassett and Pilla(35) reported on the successful use of the Faraday induction concept (PMF) to repair fibular osteotomies in beagle. Coils were placed on either side of the leg in such a manner that the magnetic fields from the coils traversed the defect and were additive (Fig. 7). The currents through each coil were pulsed in two ways, at 1 pulse s 1 and at 65 pulses s 1. As revealed by mechanical testing of the fibula subsequent to treatment, there was greater indication of recovery with 65 pulses s 1, a finding that was consistent

Figure 6. The current density J produced in tissue by a voltage source V acting through electrodes E is no different from the current density induced by a changing magnetic field B according to Faraday’s law. The B field is produced by a current I that energizes a coil C, whose plane is perpendicular to the page.

with the prediction from Eq. 2 of a larger current with higher frequency.

INVASIVE (IMPLANTED) ELECTRIC TREATMENTS

Although the use of pulsed magnetic fields provides a means by which one avoids electrode implantation, some surgeons still prefer the extra advantages that come with direct observation of the pseudarthrosis defect. In addition, there is a very lengthy literature background on delivering dc directly to bony defects.

In late 1971, groups at New York University (NYU) and at the University of Pennsylvania independently demonstrated that electrical stimulation using implanted dc devices was successful in repairing pseudoarthrosis defects in humans. In both cases, electrodes and battery were surgically implanted with provisions for percutaneously monitoring the current. Otherwise, however, the methods employed were strikingly different. The NYU group, led by L.S. Lavine (37) used platinum wire electrodes on either side of a congenital pseudarthrosis in the lower tibia of a 14 year old male, in effect allowing the current to pass through the defect (Fig. 8). The polarity of the current was such that the proximal side of the defect was negative. This approach was the same as successfully used in this group’s previous experiment (29) to repair defects in rabbit femur. The current was monitored and maintained over the 18-week treatment period at 3.9 mA.

By contrast, Friedenberg et al. (38) in treating a nonunion in the medial malleolus of a 51 year-old woman, used a technique (Fig. 8) that had been previously been found to be successful in producing callus in rabbit fibula (28,39). A stainless steel cathode was located directly in the defect

Cast

Figure 7. The PMF technique uses two flat coils connected in series to generate a magnetic field (dashed line) through the bone defect. Because the magnetic field is changing rapidly, a voltage is induced, producing a current in the vicinity of the defect. The process is completely noninvasive. In this sketch (36) the two parallel coils, whose planes are perpendicular to the page, are shown outside the cast.

and the anode, an aluminum grid, was taped to the skin. A constant-current power source maintained the current at 10 2 mA over the 9 week treatment period. Again, as with the nonunion treatment employed by the NYU group, the outcome was successful.

These differences in treatment, both leading to repair of the nonunions, remain unresolved. The one treatment (37) is consistent with prior animal work in which the proximal side of bone was found to be intrinsically negative, while the second result (38) fits those observations (24) claiming that fractures are more negative than the rest of the bone. These differences tend to highlight a key difficulty connected to the research on the electric treatment of bone. Apart from the essentially empirical nature of measurements such as the BEP profile, there is no fundamentally sound basis with which to explain the underlying mechanisms, resulting in continuing uncertainties in the clinical techniques.

A number of investigators have attempted to shed light on this question of mechanisms. Almost all such ‘‘explanations’’ have focused on the electrically related regulation of different factors: parathyroid hormone (PTH) (40), adenosine 3’, 5’- monophosphate (cAMP) (41,42), insulin-like growth factor II (IGF-II) (43), bone morphogenetic protein (BMP) (44), transforming growth factor˜-beta 1 (TGF-b1) (45), and calcium ion channel transport (46). These are contributors, in varying degrees, to the cellular signal transduction pathways controlling bone growth. However, it would be truly surprising if these factors were not involved in all types of osteogenic processes, including electrical osteogenesis. At best, such factors must be regarded as merely indicators of metabolic activity in bone. At this point in time, they provide little, if any clue as to the reason why bone is responsive to electrical stimulation.

Figure 8. Two ways of using mA-level dc currents to repair defects in bone. In one case, the electrodes are applied so as to bridge the defect. In the other case, the cathode is placed directly into the defect. Both approaches have been successful (37,38) in treating human nonunions.

Presently, there are two FDA-approved implantable direct current devices for treating bony defects, both marketed by ElectroBiology Inc. (Parsippany, NJ). These are shown in Figs. 9 and 10. The Osteogen Bone Growth Stimulator supplies 40 mA through mesh electrodes. Although the cathode is located at the defect, similar to the original placement by Friedenberg et al. (38), the current is far in excess of what was thought to lead to bone necrosis (28). Apparently, the nature of the electrodes used by Friedenberg et al. (28) may have played a role in this discrepancy. The second implantable dc device is the SpF Spinal Fusion Stimulator, prescribed for spinal fusion. The positive and negative leads, in this case carrying 60 mA, are located on either side of the repair site.

NONINVASIVE ELECTRIC TREATMENT: (CC)

One method for applying an electric current to a defect in bone in a noninvasive manner is by means of capacitive coupling (CC). The background for this technique were

experiments (47,48) in which 60 kHz sinusoidal voltages were capacitively transferred to bone cell cultures resulting in an electric field within the culture medium of 20 mV cm 1 and a current density of 300 mA cm 2. The first clinical use of this was to treat nonunions (49) (Fig. 11). An overall efficacy of 77% was achieved in a group of 22 cases with a mean time to healing of 23 weeks. The FDA-approved version of this technique is marketed by EBI (Fig. 12) As in the earlier studies on cell culture, the pictured device makes use of a 60 kHz alternating electric field that is applied to the skin on either side of the defect using disk electrodes and conducting gel. The current density within the tissue is considerably less than the levels used in cell culture, only 7 mA cm 2. The term capacitive may not be warranted for the devices pictured in Figs. 11 and 12. Unlike the lack of ohmic coupling in the earlier, in vitro studies, there is a much larger ohmic contribution to the overall impedance when disk electrodes are used. A better description for this device category might be ac (i.e., simply alternating current) instead of CC.

Figure 10. Implantable device for adjunctive treatment of spinal repair (EBI)

A wide range of parameters have been used in studying the clinical and experimental aspects of the CC signal, with various voltages applied to the skin between 1 and 10 V, and frequencies between 20 and 200 kHz. The electric field strengths generated within tissue has ranged from 1 to 100 mV cm 1 and the current densities from 0.5 to 50 mA cm 2.

NONINVASIVE ELECTROMAGNETIC DEVICES: (PMF)

The successful use of PMF (also called PEMF) by Bassett et al. (35) to repair bone defects in animals noninvasively

60 kHz ac

Cast

Figure 11. Capacitive Coupling. Electrodes are attached on either side of the bony defect external to skin (here, external to cast) supplying a 60 kHz sinusoidal signal.

Figure 12. Mode of action of EBI capacitive coupling spinal fusion stimulator (EBI).

led to a number of different devices aimed at applying pulsed magnetic fields. One such early design, successfully applied to the treatment of a tibial nonunion (50,51), made use of an iron-cored electromagnet driven by a square pulse with a repetition rate of 1 pps. However, this design suffered because the large inductive reactance of the iron core acted as a constraint on the repetition rate of the coil current pulses.

With this constraint in mind, Bassett’s group succeeded in designing (52) a low inductance air-coil system that could be pulsed at higher frequencies to repair recalcitrant pseudarthroses and nonunions in humans (Figs. 13–16). The success rate that was reported (85%) was greatly in excess of the salvage rate usually obtained by orthopedists using conventional, nonelectrical procedures. However, later (55), reviewing PMF treatments for a wider, allinclusive group of pseudarthrosis cases, including those with the worst prognosis, Bassett lowered the success rate downward, to 54%.

The pulsed magnetic field that was originally used by Bassett was (and still is) based on the saw-tooth signal common to the fly-back refresher circuit in television receivers. A saw-tooth voltage (Fig. 17) is applied to a pair of many turn coils, creating a current in both coils that generates a single magnetic field. The planes of the coils are roughly parallel, and deployed on opposite sides of the defect (see Fig. 7), creating a commonly directed magnetic field through the defect. The sawtooth signal applied to the coils results in a rapidly changing magnetic field, 10 tesla per second (T s 1), maximized at those times when the voltage applied to the coils is falling sharply. Faraday’s law results in the induced voltage shown in Fig. 18. The net induced signal that appears in the vicinity of the defect consists of bursts of 21 pulses, each individual pulse 260 ms in duration, with the bursts repeating at 15 Hz. The magnetic field that actually appears in the area of the defect rises, with each pulse, to 10 G (1 mT), before dropping precipitously in 25 ms. It is this rapid change in B that contributes the most to the induction of current (see Eq. 2). For example, if a sine wave signal of 10 G at 60 Hz were applied to the same region instead of this pulse, the maximum current would be 600 times smaller.

The various PMF clinical and experimental signal repetition rates that have been attempted vary between 1 and 100 Hz, with the maximum magnetic field intensity at the defect site ranging from 0.1 to 30 G, and the induced electric field at the site ranging between 0.01 and 10 mV cm 1.

NONINVASIVE ELECTROMAGNETIC DEVICES (ICR)

Magnetic fields are also used in bone repair in ways that have nothing to do with Faraday induction. It was shown in 1985 (56) that the results embodied in the so-called calcium efflux effect (57,58) were in close agreement with predictions based on the resonance characteristics of certain biological ions subject to the Lorentz force. Specifically, the shape of the nonlinear frequency dependence of

Figure 13. Treatment of ununited scaphoid fracture with PMF (53).

calcium binding to chick brain tissue was what might be expected for a particle with the charge-to-mass ratio of the potassium ion moving in combined parallel sinusoidal and dc magnetic fields whose ac frequency and dc intensity corresponded to the ICR condition for Kþ. This observation also explained earlier work (59,60) in cell culture demonstrating that weak low frequency magnetic fields enhance DNA synthesis in a manner that is clearly not related to Faraday induction, since the additional DNA synthesis does not scale with either frequency or intensity. Ion cyclotron resonance is a magnetic effect that is fundamentally different from Faraday’s law as expressed in Eq. 1. More specifically, as regards possible effects of magnetic fields on bone, it entails a totally different phenomenon than the induction of current in bone using pulsed magnetic fields.

Unlike previous attempts to arrive at the electromagnetic conditions required for electrical osteogenesis, exact predictions are possible using the ICR effect. One can focus on a specific ion and adjust the intensity of the dc magnetic

Figure 16. PMF device in place on patient (EBI).

field and the frequency of the ac magnetic field to ‘‘tune’’ to this ion. This is because a resonant condition occurs when the ratio of the frequency of the ac field to the intensity of the dc field is equal to the charge-to-mass ratio of the ion. The simple expression governing this resonance is

Figure 18. Voltage induced in tissue by PMF.

more likely to stimulate the gating mechanism for ion channel transport.

A great deal of work has been done in examining the effects on biological expression when tuning to Ca2þ, Mg2þ, and Kþ, not only in bone cell culture (Fig. 19) (62), but also in neural cell culture, in animal behavior, and in plants (61). It is generally agreed that ICR tuning to these ions can have striking effects on growth. One such example (63) is shown in Fig. 20 illustrating the relative effects on explanted embryonic chick femora cultured under Ca2þ and under Kþ ICR magnetic field conditions.

where v is the (angular) frequency of the ac field, in rad s 1, B is the intensity of the dc magnetic field, in Tesla, and, q/m is the mass-to-charge ratio of the ion. For practical applications, the angular frequency v is replaced by its equivalent, 2pf, where f is the frequency in hertz (Hz). The underlying interaction mechanism for this effect in living tissue is still in question (61), but the most reasonable explanation is that ions in resonance are

Signal intensity

Time

Figure 17. Sawtooth voltage applied to PMF coil.

Frequency (Hz)

Figure 19. Frequency response of insulin-like growth factor in bone cell culture under combined ac and dc magnetic field exposure 62. The dc field was maintained at 20 mT for each of the points shown. There is a clear peak at 15.3 Hz, corresponding to the predicted ICR condition for Ca2þ resonance in Table 1.

Figure 20. Effect of ICR magnetic exposures on chick embryonic growth (63). The topmost femur, the shortest, was grown under Kþ ICR tuning, while the bottom femur was grown under Ca2þ ICR magnetic field conditions. The middle femur was not exposed to any ICR field.

Diebert et al. (64) examined the efficacy of ICR in repairing defects in rabbit fibula, basically reusing the animal model that had been previously employed to study electrical osteogenesis (29), but applying an ICR magnetic field combination instead of a dc current. It was found that the 28-day ICR treatment yielded results equivalent to or better than those employing direct current and pulsed magnetic fields. For animals exposed to Ca2þ resonance magnetic fields for as little as 30 min day 1, there was an average increase in stiffness of 175% over controls, rising to nearly 300% when the exposures were maintained for 24 h. Somewhat smaller increases in stiffness were also observed for exposures tuned to the Mg2þ charge-to-mass ratio.

Another aspect of the ICR effect is that one can also use harmonics, that is, multiples of the frequency condition given in Eq. 3. For theoretical reasons (65) only odd harmonics are allowed. Thus, the most general expression for cyclotron resonance frequencies is

Figure 21. ICR bone repair device for treating nonunions. (djOrthopedics Corp.)

Corporation for treating pseudarthroses and enhancing spinal fusion (Figs. 21,22). The time variation of the magnetic field generated by these devices is shown in Fig. 23. Because the ac and dc magnetic field directions must be maintained parallel to ensure the resonance condition, these clinical devices achieve the frequency/field ratio by fixing the frequency of the applied sinusoidal magnetic field at 76.9 Hz, while using a second coil to continuously adjust for changes in the parallel component of the local dc magnetic field, to maintain this dc level at

20mT.

Some observers incorrectly use the term combined mag-

netic field (CMF), to characterize this clinical technique. It is important to understand that the fields that are combined are highly specific, following the rules expressed in Eq. 4. In addition, it is possible to achieve the same conditions in tissue with a single magnetic field, using a prepared current derived from an arbitrary waveform generator. For these reasons, the term ICR should be used for all clinical and research techniques that are otherwise termed CMF.

Table 1 lists the frequency/field ratios (fn/B) for the three ions, Mg2þ, Ca2þ, and Kþ for the first three harmonics from Eq. 4. Note that some of these ratios are numerically close to one another. The 5th harmonic of Ca2þ is slightly > 1% greater than the 3rd harmonic for Mg2þ (3.83 vs. 3.79). This observation led S.D. Smith to suggest that using a frequency/field ratio of 3.8 might be particularly effective in bone where growth is indicated for both Ca2þ and Mg2þ stimulation (66). This ratio is the basis for a number of bone stimulation devices manufactured by the djOrthopedics

Table 1.

Figure 22. ICR device for adjunctive use in spinal fusion. (djOrthopedics Corp.)

It is also sometimes incorrectly reported that ICR is an inductive procedure. However, the inductive current generated in the ICR device is negligible, approximately a factor of 10 5 smaller than the currents induced by PMF devices. While clearly noninductive, the actual ICR interaction mechanism is still in question (45). It is most likely coupled to events occurring at membrane bound ion channels (40), as evidenced that the calcium channel blocker nifedipene prevents the ICR response (67). It has been suggested, in this regard, that the channel gating process may be sensitive to the resonance tuning of specific ions (45).

SUMMARIZING EFFICACIES FOR THE VARIOUS TREATMENT

The three types of noninvasive treatments for bone defects, PMF, ICR, and CC, have each been subjected to randomized, double-blind trials and are shown to be efficacious, with an overall success rate of between 50 and 70%. One reason for this variation is undoubtedly the inclusion, or lack therein, of patients with defects that are intrinsically more difficult to repair. As the gap in a pseudarthrosis extends to widths > 5 mm, the likelihood of successful treatment diminishes. For this reason, some clinicians choose to exclude patients with radiographic gaps > 5 mm from electrical treatment (21,68).

More than 20 years after Bassett’s original use (52) of pulsed magnetic fields to repair nonunions, a definitive work on using PMF to treat delayed unions was published by Sharrard (69). A total of 45 fractures of the tibia were examined in a double-blind multicenter trial, with active PMF stimulation in 20 patients and dummy control units in 25 patients for 12 weeks at 12 h/day. The results, 9

Time

unions in the active group compared to only 3 in the control group, were ‘‘very significantly in favour of the active group

(p ¼ 0.002)’’. This effectiveness of PMF stimulation was confirmed for the case of tibial osteomoties in still another randomized, double blind study (70). Similarly, the successful use of CC and ICR, respectively, in treating nonunions, was reported by Scott and King (71) and by Longo (72). Recently, there has been increasing interest in the use of these electromagnetic techniques as an adjunctive to spine fusion. Again, as with the treatment of pseudarthroses, randomized, double-blind trials carried out for the PMF (73), CC (74) and ICR (75) techniques, have indicated that each is also efficacious in the adjunctive treatment of spine fusion.

GENERAL REMARKS

A summary of the electrical and electromagnetic treatments for nonunions and spinal fusion is given in Table 2. It is difficult to make simple comparisons based solely on the relative electrical characteristics, since each modality is based on different types of specifications, including current, current density, time rate of change of magnetic field, frequency, and magnetic intensity. As mentioned above, the levels of current that have been used to achieve osteogenesis extends over a range that has many orders of magnitude, a fact that seems to preclude any mechanism that is simply connected to current alone. It is highly likely that the larger of these successful current levels achieve osteogenesis, as Becker has suggested, by acting as an irritant, and that the smaller levels are perhaps related to the sorts of currents that might occur naturally, perhaps as the result of stress-generated potentials. One measure of

this potential dichotomy is the remarkable fact that both ICR and PMF techniques are equally successful in treating nonunions, despite the fact that the induced currents differ by a factor of 105.

There is undoubtedly room for improvement in the efficacy of the various electromagnetic treatments to repair bony nonunion. Note that both treatments, PMF and ICR, were each adopted for clinical use on the basis of the original designs, with no subsequent studies before or after FDA approval that might have been initiated to search for waveforms and signals that conceivably could be used to optimize treatment. Thus for pulsed magnetic fields, it remains to be seen what roles are played by variables such as pulse width, rise time, repetition rate, and so on, and whether marked improvements in efficacy would follow optimization of these key variables. At least one report (76) claims that peak magnetic fields 100 times smaller than used in the EBI PMF device are just as effective in treating nonunions. Similarly, positive results were obtained in treating tibial osteotomies in rabbit with very different pulse characteristics from that of the EBI clinical device (77). Not only was the magnetic pulse reduced by a factor of 15, but the pulse repetition rate was reduced by a factor of 10, and the frequency components in excess of 20 kHz were filtered from the signal. This lack of optimization is equally true for the djOrthopedics ICR therapeutic signal, based on an approximate simultaneous stimulation of Ca2þ and Mg2þions as well as a very specific ratio of ac to dc magnetic intensities. The ICR device presently approved by the FDA sets this ratio at unity, despite the fact that a number of investigators (78– 80) suggested that this ratio may have important consequences for the efficacy of the resonance interaction.

Furthermore, it has been suggested (31,81) that the fundamental reason why some electrical treatments of pseudarthroses are successful may have little to do with the nature of the electrical signal itself, but rather that the initiation of callus formation is known to be tied to local irritants, such as occurs with mechanical, thermal, or chemical sources. It is not inconceivable that the efficacy of treatments such as PMF may result from its role as an irritant. There is evidence (82,83) indicating an increased expression of heat shock proteins in response to low level electromagnetic fields. This type of genetic expression can result from a wide range of stress factors.

The fact that electrical osteogenesis occurs naturally, in growth, homeostasis, and repair and, further, that it can be brought about by exogenous application, begs the question as to whether the present 50–70% repair rate might be substantially improved with further research into the actual underlying mechanism.

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