Chapter 2

Optimization of magnetic forces for drug delivery in the inner ear

Walid Amokrane; Karim Belharet; Antoine Ferreira    HEI Centre campus, Châteauroux, France
University of Orléans, Bourges, France

Abstract

The inner ear is an important structure in the hearing system and in the balance sensibility for a human. The ear anatomy, especially the cochlea's, has a complex structure that makes disease treatment difficult. A liquid called perilymph is trapped in the cochlea by the Round Window Membrane (RWM). A leak of this liquid would result in a total loss of hearing which limits the possibility treatment by injecting drugs in the tympanic cavity, close to the RWM, to ensure that the maximum quantity of drugs is able to pass through it. The problem of this method is the side effects caused by drugs. The use of magnetic nanoparticles could be an alternative to massive drug delivery.

This chapter aims to review the current knowledge concerning the RWM structure to define the best model, close to the real structure, to push magnetic particles through it.

Keywords

RWM; Viscosity; Magnetic force; Nanoparticles

2.1 Introduction

For several years the inner ear drug delivery has become a challenge in the treatment of sensorineural hearing loss [1]. Recent advances in molecular therapy and nanotechnology have stimulated the development of a variety of methodologies such as systemic and intratympanic, or even directly intracochlear way through an injection into the inner ear [1]. The treatment of the inner ear is limited by the low concentration achieved with systemic delivery because a blood–cochlear barrier exists, which is anatomically and functionally similar to the blood–brain barrier [2]. High systemic doses must be administered in order to obtain inner ear therapeutic concentrations [3]. The necessity of long-term treatment and significant adverse effects limit furthermore the use of systemic administration [4]. It is therefore necessary to develop safe and reliable mechanisms for the direct delivery of drugs into the inner ear. Methods for local delivery can be categorized as either intratympanic or intracochlear approaches.

The minimal invasive method of local drug administration to the inner ear used for several years in clinical practice is to inject the drug into the middle ear cleft through the tympanic membrane under local anesthesia. In this way, the drug is in contact with the RWM for an undetermined time, and subsequently diffuses through the membrane into scala tympani. However, the delivered quantity of therapeutic agent to the inner ear is poorly controlled by this method since the drug is easily evacuated through the Eustachian tube to the pharynx [5]. This method is also limited by the RWM and the absence of the perilymphatic flow allowing the drug diffusion towards the apex of the cochlea. Indeed, contrary to the systemic circulation, the perilymphatic and endolymphatic flow is extremely low. Passive diffusion towards the apex of the cochlea is thus complicated due to its length, resulting in a large gradient from the base to the cochlea [6]. Side effects are lower compared to systemic delivery, but high doses are still required to achieve intracochlear therapeutic concentration. In addition, the amount of drug through the RWM is highly variable and depends on the individual anatomy [5,7]. Intracochlear injection enables precise and efficient administration of the drug. This pathway is again much less accessible and risky due to the creation of a perilymph fistula and the hydrostatic pressure generated by the injection. Because of this fistula, perilymph leakage is inevitable and the loss of injected drug must be taken into account [8].

Magnetic nanoparticles represent a promising solution as a drug carrier in the inner ear as they move into the cochlea with minimal hydrostatic force under the effect of externally applied magnetic field [9]. Magnetic nanoparticle driving was first reported using magnets [10] or superconducting magnets [11]. Martel et al. proposed to use the gradient generated by clinical MRI to pull the particles [12]. MRI driving and trajectory of magnetic nanoparticles can be modeled and preplanned [13]. These untethered particles can navigate in body fluids to allow a number of new minimally invasive therapeutic and diagnostic medical procedures. Indeed, the development of untethered microdevices that can be steered in the blood vessels could benefit several minimal invasive surgeries or interventions [14]. However, as the overall size of the micro devices is reduced, it becomes technologically more challenging to propel them. Different propulsion mechanisms have been proposed [1518]. Therapeutic magnetic microcarriers (TMMC) guided in real-time by a magnetic field gradients system are actually experienced as a means to improve drug delivery to tumor sites. One solution that has been validated in vivo is the use of the magnetic field and magnetic gradients generated by a clinical magnetic resonance imaging (MRI) system as the energy source for propulsion [18]. Magnetic iron–cobalt nanoparticles encapsulated in biodegradable poly (D,L-lactic-coglycolic acid) (PLGA) microparticles with the appropriate saturation magnetization (Ms) have been successfully used in animals. Similar wireless control navigation of intraocular microparticles has been applied in ophthalmic procedures for drug delivery purposes [19]. Complex non-uniform magnetic fields and high magnetic gradients, generated by an OctoMag electromagnetic system, allowed less invasive and safer retinal surgery and provided an increased level of dexterity desired by clinicians.

The main objective of this study is to characterize a magnetic force able to inject superparamagnetic particles through the RWM and mobilize them in the cochlea. This will enable us to optimize the magnetic actuation system for generating the magnetic fields. It is therefore important to study the anatomical and physiological characteristics of the RWM to develop a dynamic model that represents the movement of these particles in this environment.

2.2 Ear anatomy

Ear is a part of the head which has a remarkably intricate structure, able to detect balance and sound. Fig. 2.1 shows the anatomy of the human ear. It is usually divided into three parts. The first one is the outer ear [20] which includes the ear canal that is lined with hairs and glands which secrete wax. This part of the ear provides protection and channels sound. The auricle is the most visible part of the outer ear and it is what most people are referring to when they use the word “ear”.

Image
Figure 2.1 Anatomy of the human ear.

The second part is the middle ear which has two important roles: it must protect the inner ear and turn the air vibrations coming from the outer ear on structure-borne noise (analyzed by the inner ear) [21]. The middle ear consists of a chamber containing air, called the tympanic cavity, which contains a structure-borne system, tympano-ossicular chain, made up of three bones called ossicles: the hammer, anvil and stirrup. The eardrum closes the outer ear with the tympanic membrane, and marks the beginning of the inner ear which extends to the round and oval windows. In addition, it communicates with the pharynx via the Eustachian tube [22].

The last part, called the inner ear, is the deepest part in the ear anatomy. It is constituted of the vestibular system that is responsible of the balance sensibility, and the cochlea, which is dedicated to hearing. The vestibular system is a complex part of the ear, which controls the movement and position of the head. It thus allows the brain to balance the body. Control occurs through sensory cells: they react to the movement of the fluid in the inner ear, and send information to the brain as nerve impulses. Theses abilities depend on specialized receptors called hair cells.

The cochlea is a spiral shaped like a snail with a length of 31–33 mm for humans [23]. It varies in diameter along its length from apex to base (from 1 to 2 mm). Stretched across the middle of the tube is the organ of Corti, which is the highly organized basilar membrane that contains the mechano-sensory cells of the inner ear. The basilar membrane moves in response to sound waves that enter the inner ear. The structure of the organ of Corti is ton topically organized so that high frequency sounds produce the greatest motion at the base of the tube and low frequency sounds move the organ most at the apex.

The labyrinthine liquid perilymph and endolymph have a dual physiological role: they contribute to the involvement of cochlear and vestibular hair cells by transmitting the mechanical signal, and participate in the transformation of this signal into a nerve message by setting match of molecular phenomena between liquids and hair cells. While the perilymph has a composition close to other extra-cellular fluids (Na+ and Cl− near electrostatic equilibrium) [24], the endolymph is characterized by a potassium overload (K+) which results in an endolymphatic potential of +80 mV [25].

Round window is one of the two openings into the cochlea from the middle ear. It serves as a barrier between the middle ear cavity and cochlea and plays an important role in middle ear and cochlear mechanics [26]. Mechanical properties of RWM affect cochlear fluid (endolymph and perilymph) motion and thus the movement of the basilar membrane. Fig. 2.2 shows us the RWM structure. It consists of three layers from the middle ear to cochlear side: the outer layer, middle layer, and inner layer. The outer layer is divided into an epithelial layer and a subepithelial connective tissue layer which lies between the epithelial layer and the middle layer [27]. The middle layer, or core of connective tissue, contains collagen fibers, fibroblast, and other elastic fibers, and provides the main structural support for RWM. The inner layer is basically a continuation of the mesothelial cells lining the scala tympani of the cochlea. Adult human RWM is usually thicker at the edge than at the center, and its average thickness is about 70 μm [28]. Its diameter is on average between 1.81 and 2.05 mm. However, there were no mechanical properties of RWM available in the literature excepting the viscosity of each layer. Bohnke and Arnold [29] used 9.8 MPa as Young's modulus of RWM, Gan et al. [30] used 0.35 MPa, and Zhang and Gan [31] used 0.7 MPa for RWM.

Image
Figure 2.2 The round window membrane structure.

The permeability of the RWM can be influenced by the factors such as size, configuration, concentration, liposolubility and electrical charge of the substance, the thickness and the condition of the RWM. The substances placed on the RWM may traverse through the cytoplasm as pinocytotic vesicles or through different channels in between cells in the epithelium. In the connective tissue layer, cells can phagocytize the substance and traverse towards perilymph and/or penetrate blood or lymph vessels [32].

2.3 Diffusion model of magnetic particles

To demonstrate how it is possible to penetrate magnetic particles through the RWM, we start from the idea that our membrane can have several models, depending on the size of magnetic particles which must penetrate the RWM using its main parameters, namely the viscosity, the Young modulus, and the thickness from the literature. Let us define the dynamic equation of a particle as

F=Fd+Fk+Fm+W+Fvdw+Fc=mγ

Image (2.1)

where FdImage is the hydrodynamic force, FkImage the stiffness force, FmImage the magnetic force generated by an external device, WImage the weight of the particle, FwdwImage the van der Waals force, FcImage the contact forces between layer walls and the particle, m its mass, and γImage its acceleration. Fig. 2.3 shows all forces applied on the magnetic particle during its displacement. In this study, we neglect FvdwImage and FcImage. Depending on the magnetic particle size, Eq. (2.1) provides two different models of a particle's penetration through the RWM: a viscous model applied to a particle smaller than 100 nm, and a viscoelastic model applied to a particle larger than 100 nm.

Image
Figure 2.3 Forces acting on a particle.

2.3.1 Viscous model

This first model is applied to magnetic particles smaller than 100 nm, which means that we do not take into account the stiffness of the round window membrane. Like it has been defined in [34], we consider a single value of layers' viscosity, and the new dynamic model becomes

Fd+Fm+W=mγ.

Image (2.2)

Replacing the forces by their analytical expressions, we obtain

γ+6ηπrmv=V(M)Bm+g

Image (2.3)

with η being the viscosity of the cytoplasm, r, V, and v respectively the radius of the particle, volume and particle velocity, M the magnetization of the particle, B the external magnetic field, and ∇ the gradient operator. At the equilibrium (when γ=0Image) Eq. (2.3) can be written as

v=2(M)r29ηB+16ηπrg.

Image (2.4)

From Eq. (2.4) and knowing the thickness of the RWM, the injection time of the particles through the RWM is defined by

t=hRWM2(M)r29ηB+16ηπrg.

Image (2.5)

In [32] the authors show that each layer of the RWM is characterized by its own viscosity and thickness. Eq. (2.5) can be written as follows:

vi=2(M)r29ηiB+16ηiπrg

Image (2.6)

where i=[1,3]Image is the index of the RWM layer where the particle moves. We can also calculate the time required for a particle to cross each layer of the RWM:

t=i=13hi2(M)r2B9ηi+16ηiπrg.

Image (2.7)

2.3.2 Viscoelastic model

In this model, we will consider the effect of the stiffness of the inner and outer layers of the RWM. This model is applied to magnetic particles with a diameter greater than 100 nm. Using the Young's modulus as defined in [23], the stiffness of the RWM can be calculated as in [36]:

k=24EIa3

Image (2.8)

where E, I, and a are respectively the Young's modulus, inertia moment of the layer, and radius of the RWM. The dynamic model of the particle can be written as

[γv]=[6ηiπrmkm10][vx]+[VMm0][B]+[g0].

Image (2.9)

The magnetic field gradient required to propel a magnetic particle through the RWM is computed at its equilibrium state (v0Image). We can calculate the time needed to cross the RWM as

t=i=13hi2(M)r2B9ηi+g+kx6ηiπr.

Image (2.10)

2.4 Simulations and results

2.4.1 Simulation of the viscous model

In this section we will simulate the dynamics of the magnetic particle in the RWM. Specifically, we will evaluate the time needed to cross the RWM. The simulations were performed for the particles with a radius r=8 nmImage and magnetization M=1.023106 T2/AImage. The viscosity of the outer layers is the same η1,30.001 Pa sImage, and the viscosity of the middle layer is η213 Pa sImage. Fig. 2.4(A) shows the particle velocity without weight force. Fig. 2.4(B) shows the particle velocity in the case where its weight is taken into account. The minimum velocity obtained in this first case is 1.95109 m/sImage, which corresponds to an injection time of 750 min, time to cross the membrane without magnetic actuation. Applying an external magnetic force, the particle velocity increases which strongly decreases the injection time. For a magnetic force of 2.611019 NImage the injection time is 300 min, and for a magnetic force of 4.981019 NImage the injection time is 200 min, which corresponds to a gain of 57% and 71%, respectively. When the magnetic force is 5 times the weight of the particle (about 9.251019 NImage), the effect of weight is negligible as can be seen in Figs. 2.5(A) and 2.5(B).

Image
Figure 2.4 Velocity of the particle according to the magnetic force: (A) without gravity effect, (B) with gravity effect.
Image
Figure 2.5 Time of injection according to the velocity: (A) without gravity effect, (B) with the gravity effect.

Fig. 2.6 shows the velocity of a particle with a radius of 65 nm propelled with a magnetic gradient field B=0.5 mT/mImage, which corresponds to a magnetic force of 5.81016 NImage. The RWM is modeled by three layers, h1=h3=25 μmImage and h2=20 μmImage; these values are completely arbitrary because the only known information about the different layers of the RWM is that the thicknesses of the outer and inner layers are larger than the thickness of the inner layer [26,35]. This simulation shows the impact of viscosity change on the dynamics of the particle in the RWM. We clearly notice that the change in viscosity during the transition between the outer layer and inner layer results in a strong decrease of the velocity of the particle. When the particles pass through the second layer of the RWM, it is possible that they come into contact with substances, such as collagen which can significantly slow their movement. But in principle, the inner layer of the round window membrane contains essentially cytoplasm.

Image
Figure 2.6 Two different velocity levels for the outer and middle layer.

We also investigated the relationship between the magnetic force applied to a particle and its volume. Fig. 2.7 shows the evolution of the magnetic force and the particle velocity as functions of this radius. We can see that for a fixed value of the magnetic field, the magnetic force and velocity are proportional to the radius of the particle. However, from a certain radius, the particle is not able to penetrate the membrane due to its large volume. So in this case, it is assumed that the effect of the stiffness is taken into account, which increases the risk of perforation and leakage of the perilymph, and therefore a hearing loss.

Image
Figure 2.7 Magnetic force and velocity of the particle in term of its radius.

2.4.2 Simulation of the viscoelastic model

This viscoelastic model is applied to a particle having radius of 100 nm in order to introduce the elastic effect of the membrane. As shown in Fig. 2.8, the simulation results obtained with this model are contradictory to those presented in [3234]. Indeed, these results show that the nanoparticles completely crossed the RWM after a few hours under a magnetic field. The Young's modulus used to calculate the stiffness in the model gives results which indicate that the particle will never cross the round window. This could give us an indication of the parameters that we need to model the RWM, and estimate the time and the magnetic force to cross it. In addition, we can exclude this model of the round window membrane.

Image
Figure 2.8 Position of the particle obtained by the viscoelastic model.

2.5 Discussion

The first model developed aims define the dynamics of a particle in the viscous environment of the RWM cells. We noticed that the force of gravity provides a complement to the magnetic actuator, which results in a faster movement of the particle and a considerable time gain. If the viscous properties of the round window membrane are well known, we simply have to optimize our actuator to generate a sufficient magnetic force to inject a maximum of particles in the cochlea. Our goal is to inject the nanoparticles of 20 nm radius, using a magnetic gradient of about 0.5 T/m, which corresponds to a magnetic force of about 1.761017 NImage. Taking into account these assumptions, we will design the magnetic actuator able to generate this magnetic intensity. A magnetic actuator consisting of two permanent magnets able to generate both pulling and pushing force on the same axis has been proposed in the literature [35]. Compared to the electromagnetic coils, a permanent magnet system is able to generate a magnetic field of 10 to 20 times greater and a magnetic gradient 2 to 3 times higher [36].

Fig. 2.9 shows the distribution of magnetic field lines generated by this actuator obtained using the multiphysics simulation software COMSOL. The permanent magnets' dimensions are: length L=50 mmImage, width l=15 mmImage, and thickness e=15 mmImage. We can deduce that this configuration of the permanent magnets creates a node corresponding to a local minimum of the magnetic field. The presence of this node and the existence of the push force have been verified by simulation, and the results are presented in Fig. 2.9. Moreover, Fig. 2.10 shows that the magnetic gradient intersects the x-axis, which corresponds to a sign change of the magnetic force. Fig. 2.10 also shows that the magnetic gradient is higher than the previously set value (0.5 T/m), and its location relative to the device is greater than the distance between the pinna (external part of the ear) and the round window (about 3 cm).

Image
Figure 2.9 Distribution of magnetic field lines generated by the two magnets.
Image
Figure 2.10 Magnetic gradient generated by the magnetic actuator along the x-axis.

2.6 Conclusion

The injection of the nanoparticles through the round window membrane to the inner ear represents a challenge for the treatment of diseases, particularly to reach the apex of the cochlea. In the literature, it has been shown that it is possible for magnetically actuated nanoparticles to reach the cochlea, but without any indication about their localization in real time. In this chapter, two models were presented, simulated, and compared with experimental results obtained in the literature. We have shown that the dynamics of small particles depends on the physiological properties of the round window membrane, and not on the mechanical structure characterized by the Young's modulus. Once obtained, the dynamic model of the particle is able to estimate the time to cross the round window membrane. In addition, the movement of the particles depends on the magnetic gradient produced by the magnetic device. The use of the permanent magnets as actuator may be a solution because it generates both pushing and pulling force on the same axis with high intensity without a cooling system.

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