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. Author manuscript; available in PMC: 2011 Apr 1.
Published in final edited form as: Magn Reson Med. 2010 Apr;63(4):998–1006. doi: 10.1002/mrm.22268

Development of an inductively-coupled MR coil system for imaging and spectroscopic analysis of an implantable bioartificial construct at 11.1T

Nelly A Volland 1, Thomas H Mareci 1,2,4,5, Ioannis Constantinidis 1,3,5, Nicholas E Simpson 3,4,5,*
PMCID: PMC2934776  NIHMSID: NIHMS220118  PMID: 20373400

Abstract

Developing a method to non-invasively monitor tissue engineered constructs is critical for the optimization of construct design and for assessing therapeutic efficacy. For this purpose, NMR is a powerful technique that can be used to obtain both images and spectroscopic data. But the inherent sensitivity of NMR limits the observation of a bioartificial construct with current NMR surface coil technology. In this study, we address this limitation through the development of an inductively-coupled, implantable coil system, demonstrate its use at high field (11.1T), and investigate the use of this coil system for monitoring a bioartificial construct in vitro and in vivo. The results establish that large gains in signal-to-noise can be obtained with this coil system over that obtainable with a surface coil. This coil system provides a means to quantitatively analyze the structure and function of implanted bioartificial organs.

Keywords: bioartificial pancreas, inductively-coupled, implanted coil system, signal-to-noise ratio, NMR imaging and spectroscopy

INTRODUCTION

Implantable tissue-engineered bioartificial pancreatic substitutes (constructs) show potential as an alternative to current treatments of type 1 diabetes. This tissue-engineering approach has successfully restored normal blood glucose levels in animal models of diabetes for extended periods (1–4). Using NMR to non-invasively monitor an implanted pancreatic construct can provide correlations between construct function and physiologic effects post-implantation (5–7). NMR also offers the possibility of assessing changes in construct function towards developing early markers of construct failure in advance of end-point diabetic effects, e.g., hyperglycemia.

Stabler et al. developed an NMR-based method to non-invasively assess bioartificial pancreatic construct viability (8,9). Their method, correlating the 1H choline signal with cellular viability, was accomplished in vitro and in vivo on macroconstructs using a 4.7T NMR instrument and a surface coil (SC) for 1H NMR excitation and detection. However, a shortcoming of this design is the threshold sensitivity, limited to a cellular density of seven million viable cells/ml of construct under ideal in vitro conditions. Because of the desire to study implanted constructs in vivo, and to study lower cellular densities, improving the sensitivity of the NMR measurement is critical.

Two changes improve the sensitivity of NMR measurements and decrease the lower limit of detectable cells: (i) acquire NMR data at a higher field; and (ii) improve the sensitivity of the RF coil used to measure NMR. Increasing the static magnetic field strength from 4.7T to 11.1T improves the sensitivity of measurements by a factor of 2.4, since the sensitivity, characterized by the signal-to-noise ratio (SNR), increases linearly with magnetic field strength in lossy samples, such as tissue (10). Furthermore, development of an implantable coil (IC) that not only matches the size and shape of the construct, but is also inductively-coupled to an SC would increase the sensitivity of measuring specific internal structures, yet leave the subject relatively undisturbed (11–15). Additionally, these inductively-coupled implantable coil systems have an advantage over other implanted coil designs because no cable passes through the skin to connect the implanted coil to the power source and amplifiers (13,16–19). Therefore, risk of infection is lowered, increasing the period over which an animal may be safely monitored.

This report investigates the use of inductively-coupled coil systems to monitor implantable tissue-engineered pancreatic constructs. The NMR sensitivity of these systems was compared to the sensitivity of an SC at a high field (11.1T). To assess the NMR sensitivity of this coupled-coil system to previous work performed with a SC at 4.7T (8,9), comparisons were also carried out at 4.7T. Lastly, this report discusses important technical aspects toward generating inductively-coupled implantable coil systems, and issues to consider when applying these RF coils in conjunction with implanted tissue-engineered constructs.

METHODS

RF Coil development

Inductively-coupled implantable coil systems were constructed, optimized, and tested for NMR imaging and spectroscopy at 11.1T to study the NMR sensitivity dependence of RF coil design. These systems consisted of an IC inductively-coupled to an SC, functioning both as transmitter and receiver for 1H NMR (470 MHz at 11.1T). The IC surrounds a bioartificial construct, adapted from Stabler et al. (8,9), allowing the coil to be near the monitored cells (Fig. 1a). All implantable coils were loop-gap resonators, with identical inductor cross-section (1.2 cm) and thickness (2 mm) (Fig. 2c). The loop-gap resonator design was selected over a solenoid design of comparable diameter and thickness because, although a solenoid may provide a higher quality factor (20), the high inductance of the solenoid requires smaller series capacitances, and makes it harder to compensate for coating and loading effects. Since the loop-gap resonator has lower inductance, requiring larger series capacitance, it is easier to compensate for coating and loading, and is expected to perform better at higher frequency (21). All loop-gap coils were formed around a cylinder to maintain a constant cross-sectional area. A spacer verified their flatness and thickness. The inductive portion of the loop-gap resonator IC (Li on Fig. 2a) was constructed with a 2 mm-wide 202-µm-thick copper foil (McMaster Carr, Atlanta, GA). A single fixed tuning capacitor (American Technical Ceramics, Hartford, CT, CTi on Fig. 2a) was soldered directly between the copper foil ends to provide the desired resonance frequency.

Figure 1.

Figure 1

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Bioartificial construct - implantable coil assembly: (a) transverse view diagram; (b) coronal view diagram; and (c) a PDMS-coated coil used to create the assembly. The alginate beads are contained in the cylindrical cavity created by the PDMS-coated IC and the mesh screens.

Figure 2.

Figure 2

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Inductively-coupled implantable coil system: (a) circuit diagram of the implantable coil; (b) circuit diagram of the surface coil; (c) photograph of an implantable loop-gap resonator; and (d) photograph of a surface coil. Inductors (L) and capacitors (C) are shown with subscripts denoting the implantable (i), surface (s) for fixed value and (v) for variable value coils. Capacitors are similarly designated as tuning (CT) or matching (CM).

An insulating coat of polydimethylsiloxane (PDMS; Factor II, Lakeside, AZ) was added around the IC used in vitro and in vivo (Fig. 1b). To insure a uniform and sealed PDMS coating, a 2-step casting procedure was performed. A first PDMS elastomer layer was generated by curing a PDMS base for 24 hours at room temperature between glass plates carefully spaced to a desired coating thickness. Once the PDMS cured, the top glass plate was removed and several implantable coils were carefully placed on top of this PDMS layer. A second PDMS layer was then poured on top. To obtain a uniform coating thickness, a glass plate was placed atop the curing second PDMS layer and spaced from the bottom glass plate. Because PDMS is a cross-linked polymer, the two PDMS layers bond to each other, embedding the coils in a sealed coating. The coated coils were then punched out of the PDMS slab using cork borers. A smaller borer removed PDMS from the center of the coil to allow creation of the construct cavity. The PDMS coating thickness around the coils was varied to determine the coating effect on the coil resonance frequency and quality factor (Q).

The SC were constructed with a single 2 cm diameter circular loop of 2 mm wide, 35-µm-thick copper tape (Ls on Fig. 2b; 3M, Saint Paul, MN) placed on a piece of 5 mm thick Plexiglas (Fig. 2d). Figure 2b illustrates the capacitive portion including two variable matching capacitors (CMv1 and CMv2, 1–15 pf, Voltronics Corporation, Denville, NJ), one fixed tuning capacitor (CTs, American Technical Ceramics, Hartford, CT), and one variable tuning capacitor (CTv, 1–15 pf). The fixed capacitor was placed on the inductive loop, and three variable capacitors were soldered onto a circuit board to provide variable tuning and match the system to 50 Ω. A coaxial cable connected the coil to the source and preamplifier. A cable trap was added to the coaxial cable shield to block shield currents that contribute to parasitic coupling when working at high frequency (22). The SC of the inductively-coupled coil system was also used as an isolated SC because it could be tuned and matched optimally at 470 MHz. Furthermore, to compare our results with previous measurements performed at 4.7T with an SC (8,9), a separate SC was built with the same geometry as the SC of the implantable system presented above, then optimized to resonate at 200 MHz (1H frequency at 4.7T).

Coil testing

When surface and implantable coils are in proper orientation and proximity, they inductively couple. Their coupling or mutual inductance increases as the distance between the coils decreases. For surface and implantable coils tuned to the same resonant frequency, critical coupling occurs when the IC equivalent resistance equals the SC equivalent resistance within the coupled system (23). If the coupling strengthens, the coils overcouple and the resonance frequency of the system splits into two current modes: a co-rotating (+) and a counter-rotating (−) mode (Fig. 3) (24). Three different configurations of the inductively-coupled coil system were tested: 1) a loosely-coupled system where the SC was not resonant and the IC resonates at 470 MHz; 2) an overcoupled system where both surface and implantable coils resonate at the same frequency in the (+) mode); and 3) an overcoupled system where both surface and implantable coils resonate at the same frequency in the (−) mode. For the latter two configurations, the coupled coils were tuned and matched to 470 MHz. First the coil systems and configurations were simulated with antenna analysis software (Graphical Numerical Electromagnetics Code, Nittany, Inc. Riverton, UT), then constructed and optimized on the bench before testing the magnet. The characteristics of each individual coil and of the whole system (resonance frequency and Q) were determined with a network analyzer (VNA Hewlett Packard 8752C, Santa Rosa, CA) as shown on Fig. 3. The Q was measured by finding the −3 dB points (or bandwidth) from the VNA return loss plot for the SC and for the coupled-coil system, and from the VNA transmission plot for the IC (25). The VNA calibration was saved for a specific center frequency and bandwidth and checked periodically using a 50 Ω load, a short, an open for the reflection line, and a through cable for the transmission line.

Figure 3.

Figure 3

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RF coil frequency traces, simulated B1 magnetic field magnitudes, and their corresponding spin-echo images along the axis of symmetry of the coils for: a surface coil (a, d, e, & f); a coupled-coil system in the (−) mode (b, g & h); and a coupled-coil system in the (+) mode (c, i, & j) at 11.1T. All the RF coils were optimized to resonate at 470.75 MHz. In d through i, the surface coil position was set at 0 and represented by a solid line, while the IC is located at 1 cm and represented by a dashed line. (e) shows the image of a water sample when the NMR excitation is optimized for maximum signal near the surface coil. (f) shows the image of a water sample when the NMR excitation is optimized for maximum signal at the location of the tissue construct (1 cm away from the SC coil).

A sample of distilled water in a 10 mm-diameter thin-walled glass tube (Wilmad Labglass, Buena, NJ) was used to test the coupled-coil systems (with coated and uncoated IC) on the bench and in the magnet. This water phantom represents the unloaded condition because of its small size compared to the SC and the nature of the sample. To test performances under loaded conditions (similar to in vivo conditions), the implantable coated coil was placed in a mouse abdomen-like gel phantom (26–28) comprised of a viscous 6.7% (w/w) polysaccharide gel (TX-151; Oil Center Research, International LLC, Lafayette, LA). The relative permittivity of this gel phantom was determined to be 64.2 ± 1.13 (±1.77%) and its conductivity 1.13 ± 0.01 (±1.29%) S/m. Inductive coupling was accomplished by placing the water or gel phantom (with IC) on top of the SC. The IC were tested when placed around the water sample no more than a 1 cm away from the SC, and when placed in the gel phantom 0.5 cm and 1 cm from the SC.

Construct development

A biocompatible polyetheretherketone 300µm mesh (Small Parts, Inc., Miami Lakes, FL) was added on the top and bottom of the ring formed by the PDMS-coated coil to create the construct cavity (Fig. 1). The construct was acid washed to remove potential antigens and autoclaved to sterilize. Cell-free alginate beads were generated as described by Simpson et al. (29). Aliquots of freshly made beads (~0.3 ml each) were transferred into the construct.

Animal preparation and handling

Experiments were conducted with female C3H/HeN mice (20–30 g). NMR measurements and surgeries were completed under general anesthesia; induced by inhalation of 2% isoflurane in oxygen, and maintained by ventilating 1% isoflurane in oxygen. The sterile construct was implanted into the peritoneal cavity via a 2 cm midline celiotomy. The animal was placed prone onto the surface coil/cradle apparatus for NMR measurements. The construct either contained an IC, to create the inductively-coupled coil system when placed on the coil/cradle apparatus, or a coil-free construct was studied with an SC to compare an inductively-coupled implantable coil system to an SC. Animal respiration and skin temperature was monitored (Small Animal Instruments, Inc., Stony Brook, NY). Respiration was maintained between 20 and 30 breaths/min and skin temperature between 24 and 26 °C. Upon completion of NMR measurements, animals were euthanized and the construct-coil assembly or coil-free construct were retrieved.

NMR Measurements

NMR measurements were performed on an 11.1T 40 cm clear horizontal bore Magnex magnet equipped with a Bruker Avance console and on a 4.7T 33 cm horizontal bore Oxford Magnet equipped with a Bruker Biospin console (Bruker Instruments, Billerica, MA). Data acquisition and processing were done using Bruker Paravision software. The SNR was determined for each coil system with custom image analysis software written in IDL (ITT Visual Information Solutions, Boulder, CO). For coil system comparisons with phantoms, 1H images were acquired using a spin-echo (SE) pulse sequence with a repetition time (TR) of 1000 ms, an echo time (TE) of 10 ms, 1-mm slice thickness, 1 average, 4×4cm2 field-of-view (FOV) and a 256×256 matrix. PDMS signal was suppressed when required using a spectrally selective saturation pulse centered on the PDMS signal (5 ppm away from the water signal). For in vivo coil system comparisons, 1H images were acquired using a SE pulse sequence with the parameters given above, except that the TR was set to 2500 ms. Respiratory gating was used for all in vivo studies.

To compare results obtained at 11.1T and 4.7T, differences in relaxation times of the phantom water were taken into account. De Graaf et al. have shown that the longitudinal relaxation times (T1) of water and metabolites increase whereas their transverse relaxation times (T2) decrease with magnetic field strength (30). Therefore, a long TR and short TE were used at both magnetic fields to minimize signal dependence of these parameters. NMR images were acquired using an SE pulse sequence with the following parameters: TR of 15000 ms, TE of 10 ms, 1-mm slice, 2 averages, 4×4 cm2 FOV and a 64×64 matrix on both systems.

SNR measurement and statistical analysis

Three different coil-systems/configurations were tested in the magnet: 1) a surface coil (SC); 2) a coupled-coil system with an implantable loop-gap resonator in the co-rotating (+) mode; and 3) a coupled-coil system with an implantable loop-gap resonator in the counter-rotating (−) mode. The SNR of these different coil-systems/configurations were evaluated from NMR images by choosing a signal region-of-interest (ROI) at the IC position and a noise ROI outside of the sample. 770-pixel ROIs were used for each SNR determination. The Henkelman method (31) was utilized to calculate the SNR. Images in three orthogonal directions were acquired for each system/configuration. At least two independently-made coil systems were tested to assess the reproducibility of the system construction. At least two set of images were acquired for each coil system to assess the reproducibility of signal detection and the sensitivity of the systems. The SNR determined for each similar system/configuration, in the three orthogonal directions, were then averaged and normalized to the SNR measured for the SC at 4.7T (normalized to 1). This normalization was appropriate because the SC at 4.7T had the same size and shape as the SC at 11.1T. All the results are indicated as a mean ± standard deviation (relative standard deviations are indicated in parentheses in percentage). The propagation of error was included in the normalization statistics.

RESULTS

RF coil development

The loop-gap resonators and the SC were successfully built and tested on the bench and in the magnet (Fig. 2). However, the loosely-coupled coil system performed poorly on the bench because this coupled system could not be tuned or matched to 50 Ω once the IC was coupled to the SC under variable conditions of coil loading (32). Furthermore, critical coupling was never achieved in this case because the size and position of the two system coils are fixed by our experimental settings. Due to these limitations, this configuration was not considered further in this study.

The characteristics of the different coils and overcoupled systems tested are summarized in Table 1. Figure 3 shows the resonance frequencies and the B1 magnetic fields produced by a SC alone (Fig. 3a, 3d, 3e, 3f), an overcoupled system in the (+) mode (Fig. 3b, 3g, 3h), and an overcoupled system in the (−) mode (Fig. 3c, 3i, 3j). The magnetic field was calculated along the axis of symmetry of the coils, with the position of the SC set at 0 and the IC centered at 1 cm (macroconstruct location). The position of the simulation and images are aligned in the horizontal direction for ease of comparison. The (−) mode (Fig. 3g, 3h) corresponds to the higher system resonant frequency (Fig. 3b) where the currents in the two coils rotate in opposite direction and their magnetic fields are subtractive resulting in a null between the two coils (16,24). The (+) mode (Fig. 3i, 3j) corresponds to the lower system resonant frequency (Fig. 3c) where the currents in the two coils rotate in the same direction and their magnetic fields are additive. In the (+) mode, each coil of the system was tuned and matched individually to a higher frequency (490.1 MHz), whereas in the (−) mode, each coil was individually tuned and matched to a lower frequency (450.75 MHz) so that one of the overcoupled configuration resonances occurred at the desired 1H NMR resonance of 470 MHz.

Table 1.

Implantable coil, surface coil, and inductively-coupled implantable coil system characteristics: resonance frequency (f) and quality factor (Q) for surface coils at 4.7T (SC 4.7T) and at 11.1T (SC 11T), for implantable loop-gap (IC_LG−), surface coil (SC−), and coupled-coil system in the (−) mode (LG−), and IC loop-gap (IC_LG+), surface coil SC+, and coupled-coil system in the (+) mode (LG+). All coils taken individually and all systems were tested under unloaded conditions. Even when only one coil built, it was tested multiple times.

Coil # # of coils
tested		 f (MHz) Unloaded Q
IC_LG+ 7 491.81 ± 3.38

(±0.69%)		 203.90 ± 6.20

(±3.04%)
IC_LG− 4 451.4 ± 2.4

(±0.53%)		 151.8 ± 11.6

(±7.64%)
SC+ 2 493.2 72.2 ± 4.95

(± 6.86%)
SC− 1 450.8 89.3 ± 2.69

(± 3.01%)
SC 11T 4 470.75 87.0 ± 5.67

(± 6.52%)
SC 4.7T 1 200.0 69.7 ± 1.98

(± 2.84%)
LG+ 5 470.75 96.21± 3.70

(± 3.85%)
LG− 4 470.75 92.80 ± 2.65

(± 2.85%)
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Coating results

PDMS coating insulates the IC and allows it to function when implanted. Figure 4 illustrates changes in coil characteristics as the PDMS thickness around the coil is varied (n = 2 for each thickness). The coating caused the coil resonance frequency to shift (Fig. 4a, black triangles). However, the Q of a PDMS-coated coil is similar to that of a non-coated coil (Fig. 4b, black triangles). When a PDMS-coated coil is loaded (e.g., in a gel phantom), an additional frequency shift occurs (Fig. 4a, gray diamonds) and the Q decreases (Fig. 4b, gray diamonds). The frequency shift is caused by the addition of capacitance to the coil circuit whereas the change in Q results from the addition of resistance from the coating and sample. If the coating thickness altered the frequency slightly, with a maximum shift approaching 10 MHz (Fig. 4a, black squares), upon immersion, the frequency shift decreased as the coating thickness increased, with only a 3% drop for the coil with an 8-mm PDMS coating (Fig. 4a, open diamonds). Moreover, having a thicker coat reduced the coil capacitive coupling to the sample and reduced the effect of sample loading on the coil Q. The final selection of coating thickness must include consideration of these results and the practical limitation of construct fabrication (see Discussion).

Figure 4.

Figure 4

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Effects of PDMS coating thickness on the performance of the implantable coils: (a) frequency shift of an unloaded IC (▲) and a loaded IC (Inline graphic) versus coating thickness; and (b) % change (vertical scale on the right) of an unloaded IC Q (▲) and % change (vertical scale on the left) of a loaded IC Q (Inline graphic) versus coating thickness. The right-side vertical scale compares the ratio of unloaded coated coil quality factor (QUC) to unloaded, uncoated coil quality factor (QUU) and the left-side vertical scale compares the ratio of coated, loaded coil quality factor (QLC) to coated, unloaded coil quality factor (QUC). A photograph of the different coated coils tested is shown between graph a & b.

Simulations and phantom studies

Figure 3 illustrates SNR differences between the (+) and (−) modes. The ratio of B1 between the (+) and (−) modes for similar coils approaches 1.5, as shown by comparing the ratio of SNR for the loop-gap IC coupled-coil system in the (+) vs. (−) mode [1.51 ± 0.35 (± 23.1%)]. The magnetic field distribution of a SC is also represented for comparison (Fig. 3d, 3e). The non-uniformity of the image in Figure 3f is due to the overtipping of spins necessary to obtain the proper excitation at the implanted construct location. These simulations indicate that the B1 field magnitude is highest at the IC location in the (+) configuration (Fig. 3i, 3j). The ratio between the simulated B1 field magnitude at the construct location of a SC and a coupled-coil system in the (+) mode is 4.5 (Fig. 3d, 3i). This ratio agrees reasonably well with the SNR ratio of ~ 5.2 between the measured SNR at the construct location for a SC and a coupled-coil system in the (+) mode for the water and gel phantoms (see Fig. 5a, 5b).

Figure 5.

Figure 5

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SNR comparison between different RF coils under different loading conditions: (a) Unloaded coils tested with a 1-cm-diameter glass tube water sample: an SC at 4.7T (SC 4.7T) or at 11.1T (SC); coupled-coil systems with a loop-gap IC in the (−) mode (LG−), a loop-gap in the (+) mode (LG+), a PDMS-coated loop-gap in the (+) mode (LG+C), and a PDMS-coated loop-gap in the (+) mode with PDMS suppression (LG+C PDMSsup); (b) Loaded coils tested using a gel phantom: an SC at 4.7T (SC 4.7T) or at 11.1T (SC), and a coupled-coil system in the (+) mode with a PDMS-coated loop-gap placed 1 cm away from the SC (LG+C) and with PDMS suppression (LG+C PDMSsup); and (c) Loaded coils tested at 11.1T using a gel phantom: an SC, a coupled-coil system (LG+C) placed 0.5 cm away from the SC (LG+C), and with PDMS suppression (LG+C PDMSsup), and coupled-coil systems with a PDMS-coated loop-gap IC in the (+) mode implanted in vivo (LG+C in vivo) and with PDMS suppression (LG+C in vivo PDMSsup). * denotes the statistical difference between the SNR of the RF coils at 11.1T and the SC at 4.7T. # denotes the statistical difference between the SNR of the coupled-coil systems and the SC at 11.1T. Δ denotes the statistical difference between the SNR of the coupled-coil system in the (+) and (−) mode.

Figure 5a illustrates differences in SNR gain for different configurations and types of coils studied under unloaded conditions. The RF coils compared in this study include: 1) an SC at both 4.7T and 11.1T (n = 1 for each field), and 2) the two resonance modes of an over-coupled system of SC and a loop-gap IC at 11.1T (n = 7 for the (+) mode and n = 4 for the (−) mode). Of this second group, implantable loop-gap resonators (n = 4) in the (+) mode were coated with a 1-mm PDMS layer and tested inductively coupled with a SC. The SNR of SC was approximately 4 times higher at the higher magnetic field strength of 11.1T (Fig. 5a) than at 4.7T [4.07 ± 0.70 (± 17.2%), p << 0.001]. At 11.1T, all coupled-coil systems performed better than the SC. The SNR of the coupled-coil system in the (−) mode was approximately 3.4 times greater than the SC at 11.1T [14.0 ± 2.40 (± 17.1%), p << 0.001] and the coupled-coil system in the (+) mode was nearly 5.2 times greater than the SC at 11.1T [21.1 ± 3.27 (± 15.5%), p << 0.001]. The coupled-coil system in the (+) mode with a PDMS-coated loop-gap resonator IC also showed an SNR improvement of approximately 5 times that of the SC at 11.1T [20.3 ± 3.20 (± 15.8%), p << 0.001]. Furthermore, PDMS signal suppression, used with the implantable coated coils, did not significantly affect the coupled-coil system SNR [18.7 ± 2.94 (± 15.7%), p << 0.001].

Because the coupled-coil system in the (+) mode gives the highest SNR in the unloaded condition (Fig. 5a), we investigated the effect of loading on the measured SNR for a coupled-coil system in the (+) mode only. Figure 5b shows the SNR gain obtained at the construct location (approximately 1 cm away from the SC) when coupled-coil systems comprising PDMS-coated loop-gap IC (implantable-coil-construct assembly) and SC were tested under loaded conditions (n = 4). With an SC, the SNR increase at 11.1T compared to 4.7T is significant [2.38 ± 0.27 (± 11.4%), p<<0.001]. At 11.1T the coupled-coil system shows an SNR increase of more than 5.25 over the SC [12.5 ± 2.61 (± 20.9%), p << 0.001] when the implantable coil-construct assembly was placed approximately 1 cm away from the SC [0.74 ± 0.06 (8.5%) cm]. PDMS suppression did not significantly alter the results [12.5 ± 2.85 (± 22.9%), p<<0.001]. Figure 5c shows the SNR gain obtained when the distance between the implantable coil-construct assembly and the SC was approximately 0.5 cm [0.42 ± 0.05 (± 10.6%) cm in a gel phantom; and 0.62 ± 0.08 (± 12.8%) cm in vivo]. The sensitivity improvement drops to ~1.7 with the reduced distance between the two coils of the system [4.03 ± 0.90 (± 22.3%), p << 0.001 in a gel phantom; and 4.34 ± 1.51 (± 34.8 %), p << 0.001 in vivo]. At this reduced distance, PDMS suppression did not significantly alter results [4.02 ± 0.88 (± 22.4%), p<<0.001 in a gel phantom; and 4.86 ± 1.60 (± 32.9%), p << 0.001 in vivo].

In vivo studies in mice

A bioartificial construct containing a loop-gap coil was implanted into a mouse peritoneal cavity (n = 3). In vivo images (Fig. 6) were acquired within 1 hour following implantation and provided a clear visualization of the construct. The coupled-coil system shows better construct localization than the SC (n = 3), because it is more spatially selective. The average distance from the implanted construct to the mouse skin was 0.42 cm [±0.08 (±18.8%)] making separation between the IC and the SC approximately 0.6 cm, when accounting for the mouse holder. The coupled-coil system shows an SNR increase of approximately 2 over an SC at 11.1T [4.34 ± 1.51 (± 34.8 %), p << 0.001, without PDMS suppression; 4.86 ± 1.60 (± 32.9%), p << 0.001, with PDMS suppression] (Fig. 6c). With this sensitivity improvement, in vivo images of the construct allowed distinction of individual alginate beads, 400-to-700µm-diameter, entrapped within its cavity (Fig. 6b) with an in-plane resolution of ~78 µm, demonstrating the potential of this implantable coil system to directly and non-invasively monitor a bioartificial pancreas in vivo.

Figure 6.

Figure 6

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Cross-sectional image of two bioartificial constructs implanted in vivo: (a) a plain construct imaged using a surface coil; and (b) an implantable coil-construct assembly (with the same characteristic as the plain construct) imaged using the implanted coil inductively coupled to a surface coil. The white circle delineates the construct outer edge, visible in (a), but not visible when imaging with a coupled-coil system (b). An SE pulse sequence was used to acquire these images: TR = 2500 ms, TE = 10 ms, 1-mm slice thickness, 1 average, FOV 2 × 2 cm2 and matrix size 256 × 256. Respiratory gating was applied.

DISCUSSION

NMR is a powerful technique for acquiring images and spectroscopic data, but its inherent insensitivity limits application of this technique when studying implanted bioartificial constructs. To overcome this limitation, we explored the use of an inductively-coupled, implanted coil system for NMR measurements. This system has substantive advantages over an SC, and we developed a working coil system that can be implanted in vivo with the bioartificial construct. For the study of tissue-engineered constructs, this inductively-coupled, implantable coil system has the following advantages: 1) measured NMR signal is more homogeneous throughout the construct; 2) the coil system is more spatially selective without the use of selective pulse sequences because the IC surrounds the construct; and 3) if implantation depth is variable, the coupled-coil system allows tighter control over RF coil tuning and matching, and the B1 field distribution is more uniform. Additionally, when applied in vivo with respiratory gating, this system is less sensitive to breathing artifacts: the implantable coil system resonance frequency shift is small with motion, whereas with an SC the loading and sensitivity at the construct location changes more with motion.

The available SNR in the NMR measurement is of primary concern when obtaining information from implanted constructs. Increasing the static magnetic field strength improves sensitivity. Under unloaded conditions, the SNR is proportional to (B1/B2)7/4, where B1 and B2 are the magnetic field strengths being compared (B1>B2) (20), and is primarily due to the higher sensitivity obtained at higher field (33). This predicted enhancement in SNR of 4.5 between 4.7T and 11.1T is similar to our measured SNR improvement of 4.07 ± 0.18 (± 4.34%) when SC were tested. Under loaded conditions, the loading reduces the SNR gain to a linear increase with magnetic field strength (10). For field strengths considered here, the predicted SNR improvement is 2.36, agreeing well with our measured value of 2.38 ± 0.27 (± 11.4%).

The implanted coupled-coil system also demonstrates an SNR improvement over a single SC at 11.1T from a minimum of about 1.7 to over 5, depending on the implantable coil location. Coupled-coil systems in the (+) mode with a loop-gap resonator IC yield superior SNR as shown on Figure 5a, and these systems were chosen for NMR studies. Additionally, important considerations discussed previously (lower inductance and ease of construction) make the loop-gap resonator IC the best overall choice for inclusion into bioartificial constructs, particular at high fields (e.g. 11.1T).

PDMS coil coating induces mainly capacitance losses, shifting the resonance frequency, while coil loading mainly introduces resistive losses, decreasing the system Q (Fig. 4). Taking these effects into account, our results suggest that an 8-mm coating would minimize changes to IC characteristics. However, the resultant implantable coil-construct assembly is unrealistically large for monitoring constructs in vivo with research animals. The size of the mouse peritoneal cavity limits the coating thickness to no more than 1 mm. Furthermore, insertion of the coil around the construct requires the coil to be near the construct to maximize sensitivity (filling factor). Therefore, a PDMS coating thickness of 1-mm is the best compromise to maximize coil function while allowing sufficient space for a viable construct (i.e., sufficient biological component of a bioartificial construct). The coupled-coil system with the coated loop-gap IC was successfully developed, tested in a mouse abdomen-like gel phantom, implanted in a mouse, and observed in the (+) mode. Although the SNR decreases under loaded conditions compared to the unloaded conditions, there is sufficient SNR gain to obtain better images than with a SC.

With fixed coil sizes, the advantage of an inductively-coupled implanted coil system over an SC will increase as the depth of the implanted coil increases, such as in larger animal models (e.g. humans), since the SC sensitivity drops rapidly with distance (Fig. 3d, 3i). If a larger SC can obtain a greater depth of sensitivity for NMR measurements, the noise sensitivity would increase with the SC sensitive volume. Therefore, inductively-coupled implanted coils present an optimum configuration for NMR detection from implanted tissue constructs. It is important to emphasize that the approach outlined here can be applied to other situations, and is not restricted to use with implanted tissue-engineered constructs, although its inherent strength make it a desirable choice when obtaining information from implantable constructs.

CONCLUSION

Implantable, inductively-coupled loop-gap coils assembled within bioartificial constructs have been crafted, implanted into mice, and their performance analyzed. NMR images from these assemblies were easily obtained at 11.1T, demonstrating the capability of this coupled-coil system. The sensitivity improvement of this implantable coil system over SC at 11.1T is approximately 2-fold. An additional gain due to the magnetic field strength increase from 4.7T to 11.1T multiplied this improvement by 2.4. The sensitivity improvement available with the use of implantable coupled-coil systems allows significant gain in information obtained from an implanted construct, providing images with higher contrast-to-noise ratio, spectroscopy with greater SNR (e.g., 1H observation of choline for cell-containing constructs), and allowing detection of other key, but less-sensitive nuclei (e.g., 31P, 19F) critical in our efforts to quantify implanted bioartificial organ function.

With the technology of implanted coils now established, our work focuses on optimizing NMR acquisition parameters (e.g., eliminating artifacts, due to hardware and physiological issues); studying a ‘cellularized’ construct in vivo through NMR imaging and spectroscopy; and developing a multi-frequency coil system that includes important nuclei in addition to 1H.

ACKNOWLEDGEMENTS

The authors are grateful for the technical assistance of Barbara Beck, Kelly Jenkins, Garrett Astary, William Triplett, and Dr. Michelle Carman. The work described was supported by a grant from the NIH (DK47858). NMR data were obtained at the Advanced Magnetic Resonance Imaging and Spectroscopy (AMRIS) Facility in the McKnight Brain Institute of the University of Florida. Animal studies were approved by the University of Florida Institutional Animal Care and Use Committee (IACUC). The authors would also like to give special recognition to co-author Dr. Ioannis Constantinidis, who contributed greatly to this work until he suddenly passed away on April, 16th 2007.

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