Filosofía empresarial interna

Due to its high temporal and spatial resolution, MRI is widely used in the clinic and has also been a useful tool to experimentally assess LV remodeling [47, 48]. MRI exploits the permanent magnetic moment of atomic nuclei or protons. In the absence of a magnetic field, protons are naturally randomly organized, transitioning the orientation of their spin (known as precession). To obtain an MR image of an object, the object is placed in a constant (static) magnetic field B0, causing the protons to align in the

direction of the magnetic field; the protons precess about B0 at a frequency ($0)

proportional to the applied field of strength, where $0 is the Larmor frequency and % is the

gyromagnetic ratio (Equation 4.3).

(4.3)

This alignment creates a macroscopic magnetic moment that is referred to as magnetization (M). In the MR scanner, the imaged object is surrounded by a radiofrequency coil (RF coil) that is tuned to the natural resonance frequency of the protons so they can be specifically energized. The RF coil applies a series of B1 pulses

perpendicular to B0 that cause the protons to change orientation away from alignment

with B0. When the RF signal is removed, the protons realign such that their net

magnetization M is again parallel with B0, in a process referred to as relaxation. During

!

! *)!

relaxation, protons lose energy and emit a signal in the form of a free-induction decay (FID), which is detected and reconstructed to obtain 3-D MR images [47].

In order to generate a 3-D image, the FID signal must be encoded in three dimensions. While the RF pulse is running, gradient coils embedded within the bore of the MR scanner turn on to control the portion of the material that is analyzed. These gradients encode phase and frequency information within each slice, resulting in data that is generated in a series of slices to provide optimal spatial resolution and enable reconstruction of a 3-D image. Mathematical data from the MRI signal are later converted to images by a Fourier Transform for interpretation by the user [47].

The MRI scanner interface controls parameters (e.g., the type of pulse sequence and length of the pulse), which influence the magnetized protons and hence, the tissue contrast generated by MRI. Specifically, the type (e.g., spin echo, gradient echo) of pulse sequence influences proton orientation, echo time (TE) determines the time until a

signal or echo is generated, and repetition time (TR) is the time for a pulse to complete.

In combination with these parameters, MRI contrast also depends on the composition of the imaged tissue, specifically its proton density and intrinsic relaxation properties. These innate properties include the material’s longitudinal (T1) and transverse (T2)

relaxation times, which are the time required for protons to return to equilibrium or rephase, respectively, after the application of a RF pulse; differences in proton density and relaxation between different tissues differentiates them from one another. The dependence of magnetization M on these intrinsic properties and the gyromagnetic ratio is illustrated by the Bloch equation (Equation 4.4), where Mxy is the transverse

magnetization and Mz is the longitudinal magnetization [47].

! **!

(4.4)

A major advantage to MRI is its sensitivity; the RF pulse is required to influence the precession of only a few protons to successfully produce a 3-D image. However, contrast agents can also be applied in small amounts to alter the intrinsic relaxation properties (T1 and T2) of the imaged tissue or polymer [49-54]. Typical contrast agents

(e.g., gadolinium and iron oxide particles) possess paramagnetic properties that influence surrounding protons and shorten their relaxation times to enhance contrast with surrounding tissue [50-53]. Iron oxide nanoparticles are a particularly attractive contrast agent in that they are superparamagnetic (e.g., elicit stronger effects on surrounding protons) and can readily be chemically modified with functional groups, including amines and carboxylic acids [45, 49, 54]. Iron oxide nanoparticle addition leads to a shortened T2 relaxation time due to the accelerated dephasing of protons;

consequently, areas with iron oxide particles can be easily distinguished from those without [45, 49, 54]. Additionally, chemical functionalization of iron oxide nanoparticles provides a means to potentially facilitate binding with other polymers either through electrostatic or covalent interactions. This dissertation will exploit the electrostatic interactions between positively charged aminated iron oxide particles with negatively charged HA.

In addition to directly altering the proton makeup of the imaged object with the incorporation of contrast agents, imaging parameters such as the type of pulse sequence and length of pulse (e.g., TE and TR) can be adjusted. Tailoring these

properties can vary the manner by which protons reorient and, thus, can be used to

!

d"

dt

=#" $ % &

"

T

&

"

& "

T

! *+!

enhance differences in intrinsic relaxation material properties without the addition of an exogenous contrast agent [45, 55]. A non-contrast approach is attractive in that it does not add the extra variable of using a contrast agent.

This chapter will examine both x-ray and MRI as options to analyze distribution and volumes of four MeHA formulations with low or high percent modification and low or high initiator concentration (A/T) in ovine LV explants: low MeHA, low A/T and high MeHA, low A/T (used by Ifkovits et al.) and low MeHA, high A/T and high MeHA high A/T (Table 4.2) [1]. The technique that best represents hydrogel volume and distribution in explants will be used to generate data that will be input into a FE model to evaluate stress in Chapter 5.