Lipid Nanoparticle Interactions and Assembles

Novel liposome-nanoparticle assemblies (LNAs) provide a biologically inspired route for designing multifunctional bionanotheranostics. LNAs combine the benefits of lipids and liposomes to encapsulate, transport, and protect hydrophilic and hydrophobic therapeutics with functional nanoparticles. Functional nanoparticles endow LNAs with additional capabilities, including the ability to target diseases, triggered drug release, controlled therapeutic output, and diagnostic capabilities to produce a drug delivery system that can effectively and efficiently deliver therapeutics while reducing side effects. Not only could LNAs make existing drugs better, they could also provide an avenue to allow once promising non-approved drugs (rejected due to harmful side effects, inadequate pharmacokinetics, and poor efficacy) to be safely used through targeted and controlled delivery directly to the diseased site. LNAs have the potential to be stimuli responsive, delivering drugs on command by external (ultrasound, RF heating, etc.) or internal (pH, blood sugar, heart rate, etc.) stimuli. Individually, lipids and nanoparticles have been clinically approved for therapy, such as Doxil (a liposomal doxorubicin for cancer treatment), and diagnosis, such as Feridex (an iron oxide nanoparticle an MRI contrast enhancement agent for liver tumors). In order to engineer these multifunctional LNAs for theranostic applications, the interactions between nanoparticles and lipids must be better understood. This research sought to explore the formation, design, structures, characteristics, and functions of LNAs. To achieve this goal, different types of LNAs were formed, specifically magnetoliposomes, bilayer decorated LNAs (DLNAs), and lipid-coated magnetic nanoparticles (LMNPs). A fluorescent probe was embedded in the lipid bilayer of magnetoliposomes allowing the local temperature and membrane fluidity to be observed. When subjected to an electromagnetic field that heated the encapsulated iron oxide nanoparticles encapsulated in the lipid bilayer, the local temperature and membrane fluidity could be observed. DLNAs were encapsulated with different sized nanoparticles and concentrations in order to observe the effect of the bilayer nanoparticles on the lipid bilayer’s phase behavior and leakage. Two different sized nanoparticles were used, a 2 nm gold nanoparticle (GNP) much smaller than the thickness of the bilayer and a 4 nm GNP near the thickness of the lipid bilayer. The 2 nm GNPs were shown to affect the lipid bilayer differently than the 4 nm GNP. Specifically, the two nanoparticles altered the phase behavior and leakage differently in a temperature dependent fashion, demonstrating that embedded nanoparticle size can be used induce or inhibit bilayer leakage. A dual solvent exchange method was used to control the lipid surface composition of an iron oxide nanoparticle with a cationic lipid and a polyethylene glycol (PEG) lipid to produce lipid coated magnetic nanoparticles (LMNPs). PEG is well known for its ability to enhance the pharmacokinetics of nanostructures by preventing uptake by the immune system. By controlling the lipid surface composition, the surface charge and PEG conformation can be controlled which allowed the LMNPs to be used as an MRI contrast agent and a delivery system for siRNA that could be triggered with temperature.

such as Feridex (an iron oxide nanoparticle an MRI contrast enhancement agent for liver tumors).
In order to engineer these multifunctional LNAs for theranostic applications, the interactions between nanoparticles and lipids must be better understood. This research sought to explore the formation, design, structures, characteristics, and functions of LNAs. To achieve this goal, different types of LNAs were formed, specifically magnetoliposomes, bilayer decorated LNAs (DLNAs), and lipid-coated magnetic nanoparticles (LMNPs).
A fluorescent probe was embedded in the lipid bilayer of magnetoliposomes allowing the local temperature and membrane fluidity to be observed. When subjected to an electromagnetic field that heated the encapsulated iron oxide nanoparticles encapsulated in the lipid bilayer, the local temperature and membrane fluidity could be observed.
DLNAs were encapsulated with different sized nanoparticles and concentrations in order to observe the effect of the bilayer nanoparticles on the lipid bilayer's phase behavior and leakage. Two different sized nanoparticles were used, a 2 nm gold nanoparticle (GNP) much smaller than the thickness of the bilayer and a 4 nm GNP near the thickness of the lipid bilayer. The 2 nm GNPs were shown to affect the lipid bilayer differently than the 4 nm GNP. Specifically, the two nanoparticles altered the phase behavior and leakage differently in a temperature dependent fashion, demonstrating that embedded nanoparticle size can be used induce or inhibit bilayer leakage.
A dual solvent exchange method was used to control the lipid surface composition of an iron oxide nanoparticle with a cationic lipid and a polyethylene glycol (PEG) lipid to produce lipid coated magnetic nanoparticles (LMNPs). PEG is well known for its ability to enhance the pharmacokinetics of nanostructures by preventing uptake by the immune system. By controlling the lipid surface composition, the surface charge and PEG conformation can be controlled which allowed the LMNPs to be used as an MRI contrast agent and a delivery system for siRNA that could be triggered with temperature.       Targeted and controlled delivery of therapeutic agents directly to targeted tissues can be achieved, improving efficacy, lowering the necessary dose, and reducing adverse effects. Nobel Laureate Paul Ehrlich's dream of a "magic bullet" to fight disease may be realized through controlled and targeted nanoscale therapeutics (Koo et al., 2005).
In 2004, the National Cancer Institute launched the Alliance for Nanotechnology in Cancer (Alliance). The Alliance's goal is development of nanotechnology-based cancer treatments and imaging. Specifically, the Alliance is emphasizing the development of drug delivery that targets tumor cells, tumor's microenvironment, and metastatic, recurrent, and drug resistant cancers with nanotherapeutic delivery systems, theranostics, contrast agents, and complexes capable of providing multiple therapies (National Cancer Institute). The design of such multifunctional constructs is inherently complex as it requires combining different molecular, colloidal, and/or particulate agents. Furthermore, the construct must be stable, resistant to protein and immune system absorption, and capable of targeting. 4 Novel liposome-nanoparticle assemblies (LNAs) provide a biologically inspired route for designing multifunctional targeted therapeutics and imaging. The LNA structure is inspired by the early development of magnetoliposomes (liposomes with magnetic nanoparticles encapsulated in the aqueous core). Recent literature has referred to LNAs as "liposome-nanoparticle hybrids" (Al-Jamal and . LNAs are liposome structures in which nanoparticles (NPs) are encapsulated in the aqueous core, embedded in the lipid bilayer, or coupled to the bilayer surface.
Liposomes are a well-established vehicle for the administration of therapeutic and diagnostic agents (Bangham and Horne, 1964;Bangham et al., 1965;Gregoriadis, 1973; Papahadjopoulos and Ohki, 1969). As a biocompatible carrier, liposomes provide a stable means for the transportation and protection of hydrophilic and/or hydrophobic molecules. Nanoparticles are nanoscale moieties that have been demonstrated to be effective transportation vehicles, contrast agents, and agents responsive to external stimuli (such as electromagnetic fields and light). LNAs combine the advantageous properties of liposomes with functional nanoparticles to create a multifunctional therapeutic and diagnostic construct (Zhang et al., 2008).
LNAs have several advantages when utilized for drug delivery, hyperthermia, imaging, and diagnostic applications. LNAs are able to delivery hydrophobic and/or hydrophilic molecules and NPs (Zhang et al., 2009). The liposome can be modified to protect encapsulated agents from biomolecule absorption and functionalized for targeting. LNAs can also be used to concentrate encapsulates, increasing the efficiency of delivery. Also, the strategies for processing, stabilizing, and targeting liposomes are 5 well established . NPs can be magnetically guided for in vivo targeting and provide a mechanism for stimuli-responsive triggering. Surface-bound NPs also enhance the colloidal stability of LNAs and bilayer-embedded NPs can reduce spontaneous leakage Paasonen et al., 2007b;Yu et al., 2007;Zhang and Granick, 2006). LNAs harness the intrinsic advantages of a liposomal carrier, enhancing stability, bioavailability, and biocompatibility, and adds the imaging and/or responsive functionality of a NP (Zhang et al., 2008).
Drug delivering liposomes and nanoparticles have both been approved separately for clinical use by the U.S. Food and Drug Administration (FDA), see Table 1-1 and Table 1 Gregoriadis, 1973;Papahadjopoulos and Ohki, 1969). Liposomes are reliable systemic drug delivery systems because they are non-toxic, biocompatible, capable of prolonging bioavailability of encapsulated agents by reducing or preventing drug degradation and enhancing solubility and stability (Al-Jamal and .
Liposomes, as depicted in  Liposomes also open the therapeutic window, reducing adverse effects, by altering the pharmacokinetic and pharmacodynamic characteristics of the encapsulated agent (Al-Jamal and .   . This section will discuss several nanoparticles that have been utilized in LNA applications. Despite their applications in drug and gene delivery and cosmetics, cytotoxicity remains a major concern. Understanding the interactions between nanoparticles and cell membranes is crucial to NP biomedical applications and provides insight into their toxicity. NPs can be designed to bind on the cell surface, adsorb within the membrane, and translocate across the cell membrane. NPs can be exploited for novel applications by controlling the interaction between the NP and bilayer. A common way to achieve this interaction is by modifying the surface of the NP, specifically by adding positive or negative charges onto NP surface (N. Li, 2006; S. Legrand, 2008). Binding interaction between superparamagnetic iron oxide particles and stem cells are being used in cell selection process (L.F. Pavon, 2008).
NPs used in drug delivery applications can be modified to avoid drug degradation by increasing the circulation period which in turn results in cell uptake efficiency (S. Jin, 2007).

Quantum Dots
Quantum dots (QDs), 2-10nm florescent semiconductor nanocrystals, have been demonstrated as effective imaging and diagnostics agents. QDs can provide a highly sensitive contrast agent capable of exhibiting fluorescence that is 10-20 times greater than conventional imaging agents, such as organic dyes and florescent proteins. QDs are also 100 times more stable against photobleaching than organic dyes (Chan, 1998

Gold Nanoparticles
Imaging and photothermal effects of gold NPs stem from their enhanced surface plasmon resonance (SPR), where visible or near-infrared light is absorbed causing oscillation of surface electrons . SPR absorbance and the wavelength range are dependent upon nanoparticle size, core/shell configuration (e.g. silica core/gold shell (Oldenburg et al., 1999)), and geometry. Shifts in these properties are indicative of the degree of NP aggregation and/or molecular adsorption on the NP surface (Li and Gu, 2010). For photothermal therapy, absorbed light energy is converted into local heat that thermally diffuses into the surrounding medium.
Varying NP size and core/shell configuration provides a means of tuning the frequency window for photothermal therapy. It is generally accepted that gold NPmediated phototherapy is attributed to heat or resulting bubble nucleation depending on the light intensity and mode of exposure (Li and Gu, 2010

Encapsulated Liposome-Nanoparticle Assembly
Encapsulated liposome-nanoparticle assemblies (E-LNAs) are formed by encapsulating NPs within the aqueous core of liposomes ( Figure 1-2(A)). The first investigation of LNAs was inspired by the use of liposomes as a carrier for hydrophilic drugs. E-LNAs, by encapsulating NPs in the liposome core, force NPs to cluster together at a high density. High density nanoparticle loading is advantageous to hyperthermia and drug delivery because heating and drug release can be localized preventing damage to adjacent tissues. Also, high density loading provides a strong contrast agent for biomedical imaging   Shinkai et al., 1996). They can be prepared by encapsulating preformed NPs in solution or by forming NPs within the liposome core, as first shown by Papahadjopoulos in 1983(Hong et al., 1983). E-LNAs can be prepared by thin film hydration (TFH), double emulsion (DE) (Zheng et al., 1994), or reverse phase evaporation (REV) . Extrusion or sonication of post-formation liposomes can be employed to control the size of E-LNAs. Supported lipid bilayers (SLBs), NPs coated with a lipid bilayer, are formed when dcore = dNP. E-LNA formation requires the use of colloidal stable nanoparticles with a diameter (d) that is smaller than the inner diameter of the aqueous liposome core, dcore > dNP ( Figure 1-3(A)). The maximum theoretical number of encapsulated NPs is n ≈ 0.74(Vcore/VNP; V represents the volume of the core or NP), due to the close packing of 25 spheres and dcore >> dNP. Wijaya and Hamad-Schifferli demonstrated that it is possible to approach this limit, demonstrating high-density encapsulation of Fe3O4 NPs (dNP = 12.5 nm) within DPPC liposomes ( Figure 1-2(A-1)). With this design the available core volume for co-encapsulating aqueous drug molecules decreases within increasing NP concentration . However, the ability for embedding hydrophobic molecules within the bilayer is unaffected by NP concentration.
The osmotic pressure differential across the lipid bilayer and the attractive or repulsive forces between the bilayer and the NPs determine the structure of E-LNAs.
The elasticity of the bilayer determines how the LNA will deform in response to these forces. Attractive forces can include van der Waals, hydrophobic, and electrostatic interactions; and repulsive forces can include electrostatic, depletion, hydration, and steric interactions. The physical stability of a liposome-NP system can be determined

26
The adhering and non-adhering characteristics of nanoparticles can lead to changes in bilayer curvature, which can impact liposome size, shape, and phase homogeneity (Lipowsky and Dobereiner, 1998). Generally, this will occur when encapsulates are different from molecules present outside (e.g. sugars or proteins) liposomes.
LNAs are generally formed with small non-adhering NPs because NP adhesion to bilayers can significantly alter LNA structure and morphology. The exception to

Bilayer-Decorated Liposome-Nanoparticle Assembly
Bilayer decorated liposome-nanoparticle assemblies (D-LNA) are liposomes with hydrophobic nanoparticles embedded in the lipid bilayer ( Figure 1-2 nanoparticle must be less than 6.5 nm in order for the lipid bilayer to maintain its structure. Hydrophobic nanoparticles with diameters greater than 6.5 nm form micelles because they are more energetically favorable due to the high local curvature strain on the bilayer, as described in Figure 1  They showed that high NP loading with uniform distribution can be achieved in PC liposomes via thin film hydration (with sonication and extrusion). Janus particles can be prepared with embedded NPs clustered in approximately one half of the liposomes via detergent loading followed by dialysis. Clustering occurs as the liposomes minimize the energy penalty for bilayer deformationi.e. for a given concentration of embedded NPs the periodic bilayer bending energy needed to accommodate individual particles is greater than that needed to accommodate nanoparticle clusters. Park et al.

Surface-Coupled Liposome-Nanoparticle Assembly and Complexation
Surface coupled magnetoliposomes (S-LNAs) are formed when hydrophilic NPs are absorbed onto or coupled to the outer surface of the lipid bilayer ( Figure 1-2(C)). This is achieved through attractive surface interactions, notably long-range electrostatic attraction. An advantage of S-LNAs is the ease in which they can be preparedadding NPs to pre-existing liposome dispersions. Similar to bilayer embedment, decorated bilayers also provide direct heating to the bilayer in the presence of external stimuli. The design constraint for forming S-LNAs is dependent on bilayer NP adhesion and curvature. Recent investigations have shown that NPs with dNP > ~20 nm lead to the formation of SLBs due to liposome adsorption and rupture, followed by the bilayer curving around the particle ( The Granick group has shown that stable S-LNA dispersions can be formed using zwitterionic liposomes with decorated cationic or anionic NPs (< 20 nm) with a NP surface coverage above ~25% (Yu et al., 2007;Zhang and Granick, 2006). This was achieved by electrostatic attraction. Lower surface coverage led to aggregation, which demonstrates the need to balance the lipid:NP ratio. It was shown with isothermal titration calorimetry that upon binding the nanoparticles could restructure the lipid bilayer, inducing gel phases in fluid liposomes and fluid phases in gel liposomes . This observation shows that, even without external stimuli, bound NPs can induce changes in lipid phase behavior and, presumably,

Controlled Release
Controlled release of encapsulated payloads from LNAs can be induced by heating of nanoparticles raising temperature of the lipid bilayer. In vivo heating of magnetic nanoparticles has been demonstrated with external stimuli such as alternating current electromagnetic fields, microwaves, light irradiation, and lasers (Brazel, 2009). As discussed earlier, the bilayer permeability rises with temperature allowing release of encapsulates. Bothun and Preiss have demonstrated local bilayer heating due of Fe3O4 nanoparticles heated by RF causing phase transition. However, there was negligible difference between the bulk and local bilayer temperatures. Therefore, controlled release is likely due to both thermally-induced phase transition and mechanical rupture of the bilayer caused by NP heating. LNA controlled release with gold and iron oxide nanoparticle will be discussed herein (Bothun and Priess, 2011).

Gold Nanoparticles and Photothermal Effects
Heating of Au NPs is caused by the SPR properties that convert absorbed light

Iron Oxide Nanoparticles and Alternating Magnetic Fields
AC EMF operating at RF heating is due to magnetic losses being converted to heat, typically at low frequencies between 100-400 kHz. The magnetic losses for NPs < ~30 nm are due to Néel relaxation, arising from rapidly alternating magnetic dipole moments, and Brownian relaxation, arising from nanoparticle rotation and viscous losses (friction). RF heating is advantageous because it is non-invasive, easily penetrates the body, and is physiologically acceptable for up to 1 h if the product Hf, where H is the field amplitude (current  number of coils per length) and f is the frequency, is below 4.8510 5 kA/m/s

Targeted Therapy
Optimal drug delivery and biomedical imaging involves the distribution of drug and/or imaging agent to the diseased tissue while minimizing adverse side-effects to healthy tissues. Adverse side-effects limit the drug dosage that may be used during treatment, potentially requiring the dosage to be reduced, delayed, and/or discontinued. Targeting LNAs limit adverse side-effects to healthy tissues and enhance drug delivery and uptake by localizing drug delivery to specific target sites

Passive Targeting
Passive targeting of lipid-nanoparticle assemblies is facilitated by the enhanced permeability and retention (EPR) effect. Tumor growth requires a sufficient supply of oxygen and nutrients. Therefore, during angiogenesis, tumors will produce an In order for LNAs to accumulate within a tumor by the EPR effect, the residence time of LNAs in the blood must be sufficient to permeate into the tumor.
Longer circulating assemblies provide a greater opportunity to reach the tumor vascular system and enter the tumor for drug delivery and/or imaging. Nanoparticle elimination is primarily based upon the reticuloendothelial system (RES) uptake.
LNAs (as mentioned earlier) can increase their blood residence times by reducing clearance and absorption through PEGylation.
Passive targeting also exploits the unique environment created by tumors. As stated above, rapidly proliferating tumor cells require sufficient oxygen and nutrients.
However, the supply of oxygen and nutrients is typically insufficient to maintain the

Active Targeting
Active targeting involves incorporating a targeting ligand, most often antibodies, antibody fragments, vitamins, glycolipids, or peptides (Jørgensen and Nielsen, 2009), onto the surface of LNAs. In order to be effective, the targeting ligand must specifically bind to a receptor that is exclusive to tumors or is overexpressed by tumors compared to normal healthy tissue. Active targeting can be split into two Active targeting can also be accomplished with magnetic drug targeting (MDT). MDT utilizes a static magnetic field to concentration of LNAs at selected delivery sites. LNAs with encapsulated or embedded magnetic or superparamagnetic NPs can be forced to aggregate at sites with applied magnetic fields. MDT allows LNAs to be localized at an identified disease site for hyperthermia and drug delivery, minimizing the effect to adjacent tissues.   Overexpression of Her-2 frequently occurs in cervical, colon, breast, prostate, brain, bladder, and lung cancers because it allows cells to grow more rapidly. Current cancer treatments tend to lack specificity administered causing adjacent tissues to be damaged. Her-2 targeted treatment may encounter similar adverse effects to adjacent tissues because Her-2 is a naturally occurring protein. Kullberg

Magnetic Drug Targeting
Accumulation with MDT is dependent on the properties of the encapsulated magnetic nanoparticles, tumor depth, blood flow, vasculature, drug binding, and concentration. MDT is also highly dependent on the external magnetic field.

Diagnostics and Imaging
Magnetic resonance imaging is a noninvasive imaging technique that can be enhanced by the use of LNAs with superparamagnetic nanoparticles, such as iron oxide. These particles are capable of being manipulated by magnetic fields allowing them to be used as a contrast agent form magnetic resonance imaging (MRI). MRI is based upon the nuclear magnetic resonance of hydrogen protons of water in the body.
A strong magnetic field (B0) is applied to the body causing these hydrogen protons align with the magnetic field. The body is then exposed to a radio frequency (RF) pulse, transverse to B0, perturbing proton alignment with B0. Relaxation, or the realignment of protons with B0 after the RF pulse, releases energy absorbed during the RF pulse. Two separate relaxations are monitored; longitudinal relaxation, or T1recovery, and transverse relaxation, or T2-decay. T1-recovery (also referred to as "spin-lattice" relaxation) is the release of energy to adjacent tissue as hydrogen realigns with B0. T2-decay (also referred to as "spin-spin" relaxation) is the energy   This difficulty may be addressed by targeting NPs to malignant cells and tissues.
Heat transfer within tissues via NP heating can be described by a modified Pennes' bio-heat transfer model (Pennes, 1948)

Conclusion and Future Outlook
Liposomes and NPs are both well-established therapeutic and diagnostic agents. As both have been approved for clinical use, the next stage of development is to combine these two systems. LNAs combine the therapeutic advantages of these two nanotechnology systems creating a unique opportunity for achieving multi-functional therapeutic objectives. The liposomes can act to concentrate small NPs and shield them from the immune system. In turn, the NPs can be used to initiate and control drug release when exposed to external stimuli. Recent work demonstrates that there is a range of options for the design of LNAs to provide desired structures and functions.
LNAs have been demonstrated to enhance the qualities of encapsulated payloads, by providing a system capable of targeting, responding to external stimuli, and concentrating encapsulates. Also, LNAs are capable of both therapeutic and diagnostic functions.
Nanotechnology is a burgeoning new field, providing solutions to problems that were once considered unsolvable. The multi-functional quality of LNAs makes them a very exciting development in the field of nanotheranostics. However, the 61 design and use of LNAs is still in its infancy. Further investigation of the interactions between nanoparticles and the lipid bilayer is necessary to fully understand the formation, structure, and stability of LNAs. For biomedical applications, LNA bioavailability and toxicity need to be studied to grasp the full potential as multimodal nanotechnology therapeutics and diagnostics.

Areas Covered:
The current state of research and understanding of the design, characterization, and performance of LNAs. Brief reviews are provided for liposomes and nanoparticles for therapeutic application, followed by a discussion of the opportunities and challenges associated with combining the two in a single assembly to achieve controlled release via light or radiofrequency stimuli.

Expert Opinion:
LNAs offer a unique opportunity to combine the therapeutic properties of liposomes and nanoparticles. Liposomes act to concentrate small nanoparticles and 86 shield nanoparticles from the immune system while, the nanoparticle can be used to initiate and control drug release when exposed to external stimuli. These properties provide a platform to achieve nanoparticle-controlled liposomal release. LNA design and application is still in its infancy. Research concentrating on the relationships between LNA structure, function, and performance is essential for future clinical use of LNAs.  LNA release is commonly attributed to local nanoparticle heating; however, mechanically-induced release may be more plausible.
This box summarizes key points contained in the article.

Introduction
A significant challenge faced today in drug discovery is that many promising therapeutics have poor pharmacological properties, making them unsuitable for use in their native forms [1]. Some estimate that greater than 95% of new drug candidates fail to have the pharmacokinetic properties needed to be an effective treatment [2].
Improving pharmacokinetics requires chemically modifying the drug, for instance to make it water soluble, or physically modifying it by mixing or encapsulating it within a suitable matrix. Disconnect between drug discovery and drug delivery is one of the biggest reasons for the decline in breakthrough drugs in recent years [1]. New nanotechnology-based drug delivery systems have shown great potential for overcoming obstacles related to poor pharmacokinetics by providing a mechanism for controlling the delivering of low drug dosages to specific tissues or cells [3,4].
Targeted and controlled delivery can reduce the adverse effects of systemic delivery and off-target affects. The dream of Nobel Laureate Paul Ehrlich's "magic bullet" may be within reached through controlled and targeted nanoscale therapeutics.
In 2005 the National Cancer Institute provided a vision for nanotechnologybased cancer treatment that combines targeted delivery with imaging, diagnostics, and the ability to provide multiple therapies within a single nanoscale construct [5]. The design of such a multifunctional construct is inherently complex as it requires combining different molecular, colloidal, and/or particulate agents that, for example, may have different degrees of hydrophobicity or thermal instability. Furthermore, the construct must be colloidally stable, resist protein adsorption and immune system recognition, and achieve cellular targeting in its native form (i.e. without 'losing' components or cargo during circulation).
Liposome-nanoparticle assemblies (LNAs) represent a promising route for designing multifunctional therapeutic constructs. They draw inspiration from 89 magnetoliposomes (liposomes containing encapsulated magnetic nanoparticles [6-11]) and have also been referred to in recent literature as liposome-nanoparticle hybrids or liposome-nanoparticle complexes (Lip-NPs) [12,13]. LNAs consist of liposomes that contain nanoparticles encapsulated in the aqueous core, embedded in the lipid bilayer, or bound (decorated) onto the surface (Figure 2-1 Nanoparticles (NPs, up to 100 nm) have also been shown to be effective transporters, contrast agents, and agents capable of providing in vivo heating when subjected to external stimuli such as alternating current electromagnetic fields (EMFs; typically at radiofrequencies, RF) or light [14][15][16][17][18][19]. LNAs can incorporate the intrinsic properties of liposomes and NPs, providing novel multifunctional therapeutic and diagnostic vehicles. This concept was depicted by Pradhan et al. [20] for folate receptor and magnetically targeted magnetoliposomes capable of combined drug delivery and hyperthermia (Figure 2-2). Principle advantages of LNAs include the following:  Delivery of hydrophobic and hydrophilic molecules and NPs, including small NPs (< 25 nm) that are less prone to endocytic uptake due to the high curvature energy required for a membrane to 'wrap' around the particle [21].
 Strategies for processing, stabilizing, and targeting liposomes are well established [22].
 NPs can be magnetically guided for targeting in vivo and provide a triggering mechanism for controlled release (not discussed in detail herein).
The objective of this article is to provide a detailed review of LNA design and structure with an emphasis on recent work that utilize photothermal (via gold NPs) or RF heating (via iron oxide NPs) to achieve hyperthermia treatment, controlled drug release, or combined hyperthermia and drug release. LNAs containing carbon fullerenes such as C60 of C70 (i.e. fullerenosomes, see references [27][28][29][30][31][32][33][34][35][36]) are promising therapeutic structures and provide insight into LNA design, but will not be discussed herein. Likewise, NPs containing supported lipid bilayer coatings are also quite promising, but will not be discussed (see, for example [7,[37][38][39]). A discussion of reported LNA performance in vitro and in vivo will be provided. This compliments a review of "liposome-nanoparticle hybrids" by Al Jamal and Kostarelos in 2007 [12].
Recent reviews focusing on liposomes or NPs for therapeutic application, which are discussed only briefly herein, are provided in references [22,[40][41][42][43] and [4,[15][16][17][44][45][46][47][48], respectively. An expert opinion is provided that focuses on the need for more complete design principles, additional characterization of LNA structure and stability, and the validity of local heating.  The concept of a multifunctional LNA (a temperature sensitive magnetoliposome containing co-encapsulated iron oxide nanoparticles and doxorubicin) for cancer thermo-chemotherapy from Pradhan et al. [20]. Passive targeting can be achieved through the enhanced permeation and retention (EPR) effect of tumor vasculature, and active targeting can be achieved via folate receptor and by applying a permanent magnetic field. The application of an AC electromagnetic field can be used to release the drug and achieve hyperthermia treatment. Reprinted from [20] with permission.

Gold and iron oxide nanoparticles
Imaging and photothermal effects of gold NPs stem from their enhanced surface plasmon resonance (SPR), where visible or near-infrared light is absorbed causing oscillation of surface electrons [54]. SPR absorbance and the wavelength range are dependent upon nanoparticle size, core/shell configuration (e.g. silica core/gold shell [55]), and geometry. Shifts in these properties are indicative of the degree of NP aggregation and/or molecular adsorption on the NP surface [19]. For photothermal therapy, absorbed light energy is converted into local heat that thermally diffuses into the surrounding medium. Varying NP size and core/shell configuration provides a means of tuning the frequency window for photothermal therapy. It is generally accepted that gold NP-mediated phototherapy is attributed to heat or resulting bubble nucleation depending on the light intensity and mode of exposure [19]. However, recent work by Krpetic et al. [56] at low light energies suggests that photochemical effectsthe formation of free radicals during NP irradiationmay play an important role. In addition to photothermal heating, electromagnetic fields operating at RF can also be used to heat gold NPs. For example, Gannon et al. [57] examined the effect of NP concentration and RF field strength on the heating rates of 5 nm Au NPs in water. A rate of ~74 o C min -1 was measured using an 800 W RF field at a NP concentration of 67 µM.
The magnetic properties of iron oxide NPs, notably single domain superparamagnetic magnetite (γ-Fe2O3) or maghemite (Fe3O4), can also be exploited for imaging and therapy. They act as contrast agents for MR imaging, can be directed by static magnetic fields (magnetic drug delivery), and can be heated by RF (hyperthermia) [16,18,58]. RF heating is due to magnetic losses being converted to heat, typically at low frequencies between 100-400 kHz. The magnetic losses for NPs < ~30 nm are due to Néel relaxation, arising from rapidly alternating magnetic dipole moments, and Brownian relaxation, arising from nanoparticle rotation and viscous losses (friction). RF heating is advantageous because it is non-invasive, easily penetrates the body, and is physiologically acceptable for up to 1 h if the product Hf, where H is the field amplitude (current  number of coils per length) and f is the frequency, is below 4.8510 5      targeting lipids) and may provide a local heat source for both hyperthermia and drug release without adversely effecting adjacent tissue.

Core encapsulation
Encapsulating inorganic NPs within the aqueous core of liposomes is one of the simplest and earliest developed LNA configurations (e.g. magnetoliposomes or MLs [7,9]). They can be prepared by encapsulating preformed NPs in solution or by forming NPs within the liposome core as first shown by Papahadjopoulos in 1983 [67]. The later approach will not be discussed herein. Encapsualted LNAs (e-LNAs) can be prepared by thin film hydration (TFH), double emulsion (DE) [68], or reverse phase evaporation (REV) [69]. Prior to removing unencapsulated NPs or diluting, post-formation liposome processing such as membrane extrusion or sonication can be employed. The obvious design constraints are that the nanoparticles must be colloidal stable during LNA formation and that their diameter (d) must be less than that of the aqueous liposome core. When dcore = dNP these structures are referred to as supported lipid bilayers (SLBs; i.e. NPs containing a lipid bilayer coating). Based on close packing of spheres and dcore >> dNP, the maximum theoretical number of encapsulated NPs is n ≈ 0.74(Vcore/VNP) where V represents the volume of the core or NP. Wijaya and Hamad-Schifferli [70] have shown that it is possible to approach this limit, demonstrating high-density encapsulation of Fe3O4 NPs (dNP = 12.5 nm) within DPPC liposomes (Figure 2-1, A-1). With this design the available core volume for coencapsulating aqueous drug molecules decreases within increasing NP concentration.
However, the ability for embedding hydrophobic molecules within the bilayer is unaffected by NP concentration.
The structure of e-LNAs is dependent on the osmotic pressure differential across the lipid bilayer, and the attractive or repulsive forces between the bilayer and the NPs. The elasticity of the bilayer determines how the LNA will deform in response to these forces. Attractive forces can include van der Waals, hydrophobic, and electrostatic interactions; and repulsive forces can include electrostatic, depletion, hydration, and steric interactions. As classically described by Lipowsky and Döbereiner [71], adhering and non-adhering nanoparticles can lead to changes in bilayer curvature, which can impact liposome size, shape, and phase homogeneity.

Bilayer embedment
Embedding NPs into the bilayer requires that the NPs be hydrophobic and have diameters comparable to or smaller than the thickness of the lipid bilayer (~5 nm; Figure 2-1, B). LNAs formed by bilayer embedment (b-LNAs) can be advantageous as many nanoparticles are inherently hydrophobic or synthesized in organic solvents (e.g. in reverse microemulsions where the surfactant is the initial surface coating) before undergoing surface modification for aqueous environments. Similar to the ability of cells to accommodate membrane proteins, liposomes can distort to accommodate hydrophobic NPs that exceed the thickness of hydrophobic acyl region of the bilayer (~3 nm) [23,[75][76][77]. As with proteins, embedded NPs can affect lipid packing, lipid phase behavior, transbilayer permeability, and LNA structure and morphology [23,28,34,76,[78][79][80][81][82]. A unique aspect of b-LNAs (as well as surface decorated LNAs) is that the NPs can provide direct localized heating to the bilayer in the presence of external stimuli to trigger release [23,24].
It is intuitive that the size of a NP (core + surface coating) and its concentration, or more specifically the lipid:NP ratio, will influence how the lipid bilayer distorts to accommodate it and the resulting LNA structure (Figure 2-3).
Theoretical studies by Ginzburg and Balijepalli [83] and Wi et al. [84] suggest that the maximum size of a NP (dNP) that can be incorporated into a LNA while maintaining a lipid bilayer structure is ca. 6.5 nm (Figure 2-3, A and B). Above this size micellar structures are more energetically favorable due to high local curvature strain within the bilayer [84]. Experimental verification of this critical size and, furthermore, the general size effects of NPs on embedment mechanism and LNA structure are more elusive.

Surface decoration and complexation
Decorated LNAs (d-LNAs) are formed when hydrophilic NPs are absorbed onto or coupled to the outer or inner surface of the lipid bilayer ( Figure 2-1, C). This is achieved through attractive surface interactions, notably long-range electrostatic attraction. An advantage of d-LNAs is the ease in which they can be preparedadding NPs to pre-existing liposome dispersions. Similar to bilayer embedment, decorated bilayers also provide direct heating to the bilayer in the presence of external stimuli.
The design constraint for forming d-LNAs is dependent on bilayer NP adhesion and curvature. NPs with dNP > ~20 nm lead to the formation of SLBs due to liposome adsorption and rupture, followed by the bilayer curving around the particle. The critical NP diameter under which d-LNAs can be formed is dNP < 2(kb/w) 1/2 , where kb is the bilayer bending elasticity, which is dependent on lipid composition and phase state, and w is the adhesion energy.
The Granick group has shown that stable d-LNA dispersions can be formed using zwitterionic liposomes with decorated cationic or anionic NPs (< 20 nm) with a NP surface coverage above ~25% [25,26]. This was achieved by electrostatic attraction. Lower surface coverage led to aggregation, which demonstrates the need to balance the lipid:NP ratio. It was shown that upon binding the nanoparticles could restructure the lipid bilayer, inducing gel phases in fluid liposomes and fluid phases in gel liposomes [86]. This observation shows that, even without external stimuli, bound NPs can induce changes in lipid phase behavior and, presumably, permeability.  (Figure 2-1, D1 and D2) or the NPs bind to multiple liposomes and act as "bridges" (Figure 2-1, D2). Voldokin et al. [13] have shown that either structure can be formed from the same anionic Au NP-cationic liposome by used Fe3O4 NPs coated with histidine groups to bind to and complex zwitterionic/cholesterol liposomes containing Cu(iminodiacetate)-functionalized lipid.
The objective was to demonstrate a potential method using histidine-Cu(IDA) binding to form c-LNAs, thereby concentrating a therapeutic and an imaging agents at a target site. The resulting aggregates ranged from 20-100 µm in diameter.

LNA Controlled Release
This section reviews recent work on gold or iron oxide NP-mediated release from LNAs. Articles that apply these principles in vitro or in vivo are presented in section 4.

Gold nanoparticles and photothermal effects
Utilizing the photothermal heating of Au NPs, Paasonen et al. [24] demonstrated the ability to control the release of calcein (622. 6

Cellular uptake and drug delivery
Chithrani et al. [94] -19) showed that the b-LNAs were internalized by endocytosis and localized in endosomes. Exposure to UV light at 400 mW/cm 2 for 300+ s led to calcien release from b-LNAs, but not from liposomes that did not contain embedded nanoparticles.

Drug delivery and hyperthermia
Pradhan et al. [20]    Mechanically-induced release where the bilayer is 'broken' due to NP rotation or oscillation at or within a LNA bilayer.
 What is LNA toxicity and how does the design strategy affect? It is unclear if LNA toxicity will stem from the liposomes or the NPs, or if a synergistic effect will occur. Toxicity, which is important from a clinical perspective, will depend on the LNA design and associated colloidal stability. For example, toxicity could be 'low' if an LNA can retain its NP agent until it reaches a target site. This could be the case for bound NPs formed by embedment or encapsulation. However, 'higher' toxicity could occur if the NPs are released during circulation (i.e. toxicity of liposomes + NPs). Release of encapsulated NPs could be attributed to liposome fusion or bilayer disintegration, release of embedded NPs may occur due to bilayer solubilization by surface-active agents, and release of decorated NPs may occur due to charge screening or competitive binding.
 What clinical challenges exist to LNA-based therapeutic assemblies? As with toxicity, it is unclear if LNAs present unique clinical challenges beyond those reported for liposomes or nanoparticles [3,4,22]. These challenges include achieving biocompatibility, bioavailability, and cellular targeting and uptake. LNA structure, function, and stability will clearly impact how these challenges are addressed.

Declaration of interest
This work was supported by grants from the National Science Foundation (CBET-0931875) and the NASA Rhode Island Space Grant Consortium.

Introduction
Magnetoliposomes (MLs) consist of iron oxide nanoparticles encapsulated within lipid bilayer vesicles or liposomes [1][2][3][4]. Two common magnetoliposome structures formed by encapsulating hydrophilic magnetic nanoparticles are shown in can be achieved physically, by guiding the MLs using magnets placed on a body's exterior [5], as well as chemically using cationic lipids or ligand-conjugated lipids with selectivity for specific cellular receptors [5,6]. Pradhan et al, for example, recently showed that the inclusion of a folate-conjugated lipid can be used to target MLs to cancer cells via folate receptor (FR) binding and FR-mediated endocytosis [7].
A unique property of MLs is the ability to heat the encapsulated magnetic nanoparticles, typically iron oxide (maghemite, -Fe2O3 or magnetite, Fe3O4), using external alternating current electromagnetic fields (EMFs) operating at radiofrequencies (RFs) [6,8]. With this mode of heating, magnetic losses are converted to energy. The magnetic losses are attributed to Neel relaxation, which is due to rapidly alternating magnetic dipole moments, and Brownian relaxation, which is due to nanoparticle rotation (i.e. friction losses) [9]. ML heating using physiologically compatible or benign RF strengths can provide local hyperthermia treatment [10] or combined hyperthermia and drug delivery [7]. This is attributed to radiofrequencies easily penetrating the body and being non-invasive for up to 1 hour if 137 the product Hf, where H is the field amplitude (current  number of coils per length) and f is the frequency, is below 4.8510 5 kA m -1 s -1 [11]. With respect to drug delivery, initiating and controlling drug release from MLs has been attributed to the ability to manipulate the phase behavior and diffusivity of the lipid bilayer [7,[12][13][14].
Combining the ability to guide MLs using magnets and trigger release with RFs can overcome a major challenge of 'conventional' liposomes for drug deliveryobtaining high liposome accumulation at a target site and achieving a desired controlled release profile [15].

138
Challenges to ML-based hyperthermia and drug release include measuring changes in temperature without the use of invasive probes and at the site of interesti.e. local temperatures in the vicinity of the nanoparticles [8]. With respect to hyperthermia, the local temperature in vivo is important in cancer treatment where elevated temperatures must be maintained for a period of time to cause cell death (e.g. 42 o C or a ΔT of 5 o C from physiological temperature for >30 min) [16]. In this case questions arise concerning what minimum cellular nanoparticle (or ML) loading is needed to achieve sufficient heating, as well as the ability to selectively heat the cancer cells or tumor mass and not damage healthy cells or tissues [9]. In turn, for drug release applications, the heat delivered to or generated within the MLs may be used to control transbilayer release by raising the bilayer to its melting temperature [1,12,17]. However, there is some controversy associated with this statement. Keblinski et al. [18] have shown that the theoretical temperature difference between a nanoparticle surface and the bulk phase due to EMF heating is almost negligible. This was further verified by Gupta et al. [19] for Fe3O4 nanoparticles heated by RFs. In this work quantum dots were mixed or covalently anchored to the nanoparticles as temperature probes to compare bulk verse local temperature.
In this work we describe an in situ spectrofluorometric-based approach to determine the effect of RF heating on the temperature of liposomal bilayers in a Fe3O4 nanoparticle-liposome dispersion containing MLs. The approach is based on the anisotropy, <r>, of a lipid bilayer probe molecule, diphenylhexatriene (DPH), and the proportionality among anisotropy, bilayer viscosity, and lipid phase state, which depend on temperature [20]. Nanoparticle-liposome dispersions were prepared with dipalmitoylphosphatidylcholine (DPPC) and contained 25 mol% cholesterol and 0.2 mol% DPH. While cholesterol is known to stabilize liposomes, in this work it was added specifically to broaden the main phase transition or melting temperature region [21,22]. This provided a large heating window for anisotropy measurements.
Calculated (via <r>) and measured temperatures were compared to differentiate between local and bulk heating relative to MLs as a function of time during on/off RF cycles.

Chemicals
Dipalmitoylphosphatidylcholine (DPPC) and cholesterol were purchased from Avanti Polar Lipids, and diphenylhexatriene (DPH) from Sigma Chemical Company.
The aqueous Fe3O4 nanoparticle solution was purchased from Ferrotec GmbH (EMG 705 ferrofluid, 3.9 vol%). Previous work has shown that in this solution the particles are well dispersed with an average diameter of 12.5  3.4 nm [23]. Sterile deionized water was obtained from a Millipore Direct-3Q purification system. All materials were used as received with the exception of dilution.

Preparation of Fe 3 O 4 nanoparticle-liposome dispersions
Samples were prepared by reverse phase evaporation [24] (Buchi Rotavapor R-215, Zurich, Switzerland), similar to the procedure by Wijaya and Hamad-Schifferli Thermo Electron X Series, Waltham, MA). Encapsulation was calculated based on the 57Fe isotope.

Cryogenic transmission electron microscopy (cryo-TEM)
The nanoparticle-liposome dispersions were prepared for cryo-TEM at

Cryo-TEM and DLS analysis
Cryo-TEM micrographs of the nanoparticle-liposome dispersion taken one week after preparation are shown in Figure 3 It should be noted that the percentage of liposomes containing nanoparticles or nanoparticle aggregates is not equivalent to encapsulation efficiency. An encapsulation efficiency of 67% was measured one day after sample preparation based on ICP-MS analysis of precipitate and supernatant fractions after centrifugation [10].
Comparing cryo-TEM and encapsulation efficiency results suggests that most of the nanoparticles were encapsulated as aggregates within a small population of liposomes and/or that encapsulated nanoparticles were released due to ML rupture. Nanoparticle release from MLs has been observed by Wijaya and Hamad-Schifferli for high-density nanoparticle-loaded vesicles (HNLVs) prepared using DPPC and the same ferrofluid [23]. In their study the samples were analyzed 30 min after preparation and release was attributed to incomplete closure, which reduces encapsulation efficiency. In the present work, cryo-TEM was conducted one week after ML formation (stored at 25 o C) and rupture was still clearly evident.  (5) nanoparticle aggregates. The scale bar represents 500 nm and is common to both images. 3) for determining temperature as a function of DPH anisotropy during RF-heating.

Figure 3-4. Melting transition determined by DPH anisotropy.
The reduction in DPH anisotropy with increasing temperature reflects an increase in bilayer fluidity, or decrease in microviscosity, as the bilayers transition from a gel to a fluid phase.
DPH anisotropy was measured in situ during RF-heating (5×10 5 kA m -1 s -1 ) as a function of time. Heating was initiated at time zero and cycled between off and on.
The T versus <r> calibration curve (Eq. 3) was used to calculate temperature. Prior to conducting the experiment, a cuvette containing water was heated by RF for 1 h. The temperature rise from this test was from 25 to 26 o C, which demonstrates that heating of the cuvette, cuvette holder, and base within the spectrophotometer sample compartment was minimal and its contribution to sample heating was negligible.

147
A rapid increase in the calculated temperature (via <r>) was observed with EMF heating up to 1400 s (Fig. 3-5). The temperature changes with on/off EMF cycling, between 1400 and 1450 s and 1450 and 2100 s, demonstrate the response and reversibility of ML membrane fluidity and its use for remote temperature measurements. The correlation between calculated and measured (via probe) temperatures suggests that the calculated temperatures reflect that of the bulk as opposed to local to the ML bilayer. Hence, bilayer temperature could not be differentiated from bulk temperature. This finding was initially counterintuitive considering that heat originated within the MLs before being transferred through the bilayer into the bulk aqueous phase. To resolve this, a simple heat transfer analysis was adopted from Keblinski et al. [18], where they examined nanoscale versus global  DPH anisotropy (<r>, black squares) was measured at a RF field strength of 5×10 5 kA m -1 s -1 . The field was cycled randomly between off and on. Calculated temperature (red triangles) was determined from anisotropy using Eq. 1. The results were fitted by a fourth order spline fit with an exponential correlation function, and the fitted line is intended to guide the eye and depict general trends. Temperature measurements of the bulk phase (red squares) are shown for comparison.

Conclusions
The fluorescence-based approach provides a means of measuring the lipid bilayer temperature in nanoparticle-liposome dispersions subjected to EMF heating. While the intent of this work was to provide proof-of-concept using Fe3O4 nanoparticles, the technique is applicable to other inorganic nanoparticles provided they do not interfere with the spectroscopic measurements. Based on the experimental results and a simple heat transfer analysis, the temperature of the bilayer in the MLs employed was equal 149 to that of the bulk aqueous phase. This was due to rapid heat dissipation from the nanoparticle surface where the difference between the surface and bulk temperatures was negligible. We emphasis that in this work the MLs were not subjected to postformation processing, such as sonication, extrusion, or purification. Hence, the MLs were heterogeneous and contained unencapsulated iron oxide nanoparticles and nanoparticle aggregates. However, based on the heat transfer analysis and inferred by previous theoretical work [18], liposomal heating is not dependent on the ratio of encapsulated to unencapsulated nanoparticles, but rather on the total nanoparticle concentration within the dispersion. This suggests that nanoparticle-lipid interactions.

Introduction
Liposomes are well-established biocompatible carriers capable of protecting, transporting, and delivering hydrophobic cargo (in the bilayer) and/or hydrophilic cargo (in the aqueous core) for biomedical applications. 1, 2 Liposomal drug delivery systems can increase therapeutic effectiveness, increase stability, target diseased sites, and control release while reducing overall toxicity and side effects 3

Gold Nanoparticle Synthesis.
Dodecanethiol-stabilized gold nanoparticles (DDT-GNPs) were synthesized via an arrested precipitation method previous described for silver 29 , then modified for gold. 16 In short, 330 mg of gold chloride trihydrate (Acros Organics, 99%) was dissolved in 20 mL of DI water and 6 g of tetraoctylammonium bromide, TOAB, imaging. This is depicted in Figure 4-3F1-F2 where a D-LNA with a 'dark spot' is observed, but only when the TEM is over-focused do we see that this spot is comprised of a GNP cluster. This does not mean that all structures in the D-LNA samples contained embedded GNPs -it is likely that there were also 'empty' liposomes present. Additional TEM work is needed to confirm these assertions.      Table 4-2. The The pretransition and melting transition appear as peaks in the excess heat capacity (Cp, Figure 4-6). 33,34,36 8,15,20,41,42 (ii) changes in bilayer structure and mechanics such as thickness and elasticity, respectively; 9,[15][16][17]19 ; and (iii) changes in liposome size and structure, (also inferred by this and previous work) and the formation of non-bilayer structures. 8 These mechanisms are interrelated and likely occur simultaneously, necessitating additional work to determine the how these mechanisms contribute to the release behavior.

A B
Additional insight can be gained by considering molecules that partition into lipid bilayers and are known to influence lipid organization, which in turn influences inter-lipid interactions that drive phase transitions and the kinetics associated with these transitions. The characteristic time of a pretransition has been reported as 5 ± 2 min. 43 With a DSC scan rate of 1 o C min -1 , results for DPPC show that the pretransition peak present over a 3 o C temperature range is consistent with this characteristic time. In addition to GNPs disrupting lipid organization, either locally or globally within the bilayer, the GNPs would also exhibit low translational diffusion times within the bilayers that would impact the kinetics associated with lipid reorganization. This was apparent during the D-LNA pretransitions, where they were suppressed or merged with the melting transition at higher temperatures, and the greatest for GNP3-4 which occupy more space within the bilayer and are less mobile than GNP2. Given that CF leakage is greatest at the interface between phase domains, the net effect of inhibiting or delaying a phase transition would be lower CF leakage or an apparent temperature lag.

Conclusions
As shown for the GNP D-LNAs, the permeability and phase behavior of liposomes can be manipulated by the size and concentration of embedded nanoparticles at low volume fractions within the bilayers. A lipid bilayer is approximately 4-5 nm thick, and nanoparticles that were smaller than the bilayer thickness (GNP2) affected the bilayer differently than nanoparticles that were closer to the bilayer thickness (GNP4 Polyethylene glycol (PEG) is a hydrophilic biocompatible polymer commonly used to improve the blood circulation half-life, surface hydrophilicity, and reduce cytotoxicity of nanoparticles. [2][3][4][5] PEG is the most heavily studied surface modifying polymer for improving the efficiency and effectiveness of therapeutics. PEG has the ability to make a particle "stealth", improving its pharmacokinetics by reducing mononuclear phagocyte (immune) system uptake. 6 However, there is no consensus as to what the optimum PEG coverage-density, conformation, and molecular weight to prevent necessary to prevent uptake by the immune system. 3    The effect of the surface composition of PEG on the LMNPs was assessed by measuring the hydrodynamic diameter, DH, and zeta-potential, ζ, using dynamic light scattering (DLS). The DH of the SPION core in chloroform was measured to be 28.3 ± 6.8 nm. The number-weighted mean DH as a function of mol% PEG are shown in where a is the PEG monomer length and N is the number of repeat monomer units (parameters for calculations are in Table 5-1). 3 The mean distance between PEG groups, D, is calculated with the following equation: 191 where APEG is the area that a PEG chain occupies. APEG is calculated with the following: where SNP is the surface area of the iron oxide nanoparticle (2516 nm 2 ) and NPEG is the number of PEG on each nanoparticle. NPEG is calculated based upon the mol% PEG2000 and the cross-sectional area of the acyl lipid tails, ALIPID, of DOTAP and DMPE (the PEG anchor lipid). 3 PEG takes on the mushroom conformation at low PEG density, D > RF and RF ≥ L. 2,15 In the mushroom regime, PEG chains have a random coil conformation with no interaction with adjacent PEG chains. When D ≈ RF, individual PEG chains begin to take on the brush conformation. The brush conformation occurs when PEG chains are close enough to have some interaction that disrupting the ability of PEG chains to retain their mushroom conformation. 15 PEG is in the brush conformation when RF < D < 2RF. The dense brush conformation occurs at very high PEG densities when there is an increased amount of inter-chain interactions and steric repulsion forcing the PEG chains further away from the nanoparticle surface. The dense brush conformation occurs when D < RF and L < 2RF. 3, 15 PEG2000 should be in the mushroom conformation below 9 mol% PEG, in the brush conformation from 9 to 70 mol% PEG, and in the dense brush conformation at greater than 70 mol% PEG (  The size of the LMNPs was calculated with the following equation: where L is the length of the PEG chain, DNP is the diameter of the iron oxide nanoparticle core (28.3 ± 6.8 nm), and TB is the thickness of the lipid bilayer. TB for DOTAP and DMPE (the anchor lipid for PEG) is 4.2 nm 16 and 3.6 nm 17 , respectively.
The calculated L, FP, and DLMNP and the measured DLMNP are recorded in Table 5-2. At 0 mol% PEG, the LMNPs are completely covered with DOTAP. Therefore, L = 0 nm and TB was the thickness of DOTAP. The calculated and measured DLMNP were 32.5 nm and 31.7 ± 12 nm, respectively.
PEG2000 should take on the mushroom conformation below 9 mol% PEG.
When PEG2000 is in the mushroom conformation, LMushroom = RF = 3.4 nm. 3 PEG2000 should take on the mushroom conformation below 9 mol% PEG. Only the 5 mol% PEG LMNP sample should be in the mushroom conformation. The calculated DLMNP = 38.7 nm. The calculated size does fall within the standard deviation of the measured size at DH 29.9 ± 11.1 nm.

193
The measured size may be smaller than expected because the PEG chains are methoxy-terminated giving them an anionic charge. When polymer chains are not repelled from the nanoparticle surface, a special "pancake" conformation has been observed. 18 PEG lies along its own length on the nanoparticle surface. The PEG length in the pancake conformation, Lpancake, is calculated as: where, N is the number of repeat monomer units (44). 3,18,19 For PEG2000, Lpancake is 0.2 nm. The PEG2000 lipid anchor is DMPE which has a bilayer thickness of 3.6 nm. 17 Therefore, the calculated size 5 mol% PEG LMNPs is 32.
where L is the position of the shear plans, A is the surface potential, and κ is the inverse Debye length. 3,22 Therefore, as the PEG surface density and/or distance away from the surface of the charged particle increases, the LMNP ζ decreases. [23][24][25][26] Surface charge can be tuned by changing the LMNP lipid surface composition, as can be seen in brushes on block copolymer micelles. 28 The ability to control the surface composition with the DSE method allows the benefits of PEGylation to be applied to SPIONs to form multifunctional LMNPs.
Along with PEG's ability to improve circulation half-life, biodistribution, efficacy, and safety, it also has been shown to improve the r2 relaxivity allowing it to be used as a negative MRI contrast agent. Increasing concentrations of surface PEG has been show to increase the r2 relaxivitiy by thickening the hydration shell around the LMNPs. 9, 10 The effect of surface PEG concentration on r2 relaxivity of LMNPs is displayed in Figure 5-3(a). concentrations, the ζ is more cationic and there are less PEG chains to obstruct siRNA from bindng to the LMNP surface. Therefore, more siRNA could attach to the surface at lower mol% PEG, but it attachment does not strictly follow ζ.
The amount of siRNA bound to the LMNP surface was dependent on PEG conformation and temperature, as shown in Figure 5-3(b). In the mushroom conformation (0-15 mol% PEG), more siRNA was able to bind to the LMNP, due to the hightened ζ. PEG becomes more hydrophobic with increasing temperature causing the PEG to collapse and potentially interdigitate with the lipid monolayer. 30 Therefore, in the mushroom conformation, as the temperature rises more cationic charges may be available for siRNA to bind to the LMNP due to the change in PEG as temperature increases. For the brush conformation (>15 mol% PEG), the percent of bound siRNA also decreases with increasing temperature. Under these conditions, the temperature causes the PEG to become more hydrophobic, expelling siRNA from the surface.
These results indicate that LMNPs can be used for the controlled release of siRNA.
LMNPs also can generate heat when subjected to alternating current electromagnetic field operating at radio frequency (RF), allowing them to be used for hyperthermia and/or triggered siRNA release. RF heating was performed using a 1 kW Hotshot™ set to 300 A (actual output ~227-235A and 362 kHz) for 30 min. The change in temperature from the recorded room temperature was recorded and is 201 displayed in Figure 4 (a). The heating capacity of nanoparticles in an RF is measured by the specific absorption rate (SAR): where cp is the average heat capacity of the sample, in this instance water (C = 4.18 J/g K), mV is the mass of the sample volume, and mSPION is the mass of the SPIONs.
∆ ∆ was calculated based on the initial slope of the heating curve. SAR for 15 and 50 mol% PEG LMNPs were subjected were 52.8 W g -1 and 11.2 W g -1 , respectively. This ability to heat when subjected to RF The percent of siRNA released from the LMNP surface when subjected to RF heating is shown in Figure 5-4

MRI Sample Preparation
LMNPs were suspended in 1% agarose at concentrations of 1, 10, and 20 µg/mL Fe3O4 in 4 mL plastic sample tubes. 1% agarose gels were prepared by mixing agarose with 10X TBE buffer (diluted to 1X) in a flask on a hot plate set to keep the gel at 80°C until the agarose is completely dissolved. The gel and LMNP sample was then pipetted into a sample holder, vortexed, and stored at 4°-8°C until imaged.

MRI Methods for r 2 Relaxation
Samples were imaged using a Siemens Prisma 3T scanner. The sample holder was placed in a 64 channel head receive array. LMNP samples in 4ml vials were scanned using spin echo (r2). Cross section images of the vials were obtained with voxel size of 0.78mm and slice thickness of 4mm. Repetition time was 2400ms for all sequences. For the spin echo acquisition 24 echoes were collected over the range of 9 -216ms (9ms step). The inversion recovery data were taken with inversion times of 100, 200, 300, 400, 600, 1000, and 1500ms. Relaxation time constants were determined using a nonlinear least squares fit for pixel intensity vs echo time for r2.
Three-parameter nonlinear least squares fitting routines (M0, T1,2. DC offset) were used for r2 to take into account through-slice dephasing effects 31 . Relaxivity was calculated as a linear fit of relaxation rates to iron concentration.