“I want to build a billion tiny factories, models of each other, which are manufacturing simultaneously. . . The principles of physics, as far as I can see, do not speak against the possibility of maneuvering things atom by atom. It is not an attempt to violate any laws; it is something, in principle, that can be done; but in practice, it has not been done because we are too big.” – Richard Feynman, Nobel Prize winner in physics
“Nanomedical physicians of the early 21st century will still make good use of the body’s natural healing powers and homeostatic mechanisms, because, all else equal, those interventions are best that intervene least.”- Foresight Nanotech Institute
What is Nanomedicine?
Nanotechnology is a multidisciplinary field which includes the combination of such fields in science as: applied physics, materials science, interface and colloid science, device physics, supramolecular chemistry, chemical engineering, mechanical engineering, and electrical engineering. Nanotechnology can be defined as the design, synthesis, characterization and application of materials and devices with their smallest functional organization, at least in one dimension, being on the nanometer scale. Nanomedicine is the application of nanotechnology to the medical field. Because of the extremely small size of nanostructures and nanodevices, they can interact with human biological systems at a subcellular or molecular level. One nanometer (nm) is equal to one billionth of a meter. To put a nanometer into prospective, it is the approximate width of six carbon atoms or ten water molecules. A red blood cell has an approximate diameter of 7000 nm, while that of a strand of hair is approximately 80,000 nm.
What if the very first cancer cell could be detected and removed before it ever developed into full-blown cancer? What if a broken component of a cell could be removed and repaired, or replaced with a miniature biological machine? What if small pumps, the size of molecules, could be implanted into the body and deliver life-saving drugs precisely when and where needed? Nanomedicine is the area of research in which hopefully this will all someday be possible. Nanomedicine is an extremely broad field of research and is still very new. Presently, the most advanced and successful areas of research in nanomedicine are in drug and gene delivery, and imaging. The more long-term goal of nanomedicine is the development of programmable and controllable microscale robots made from nanoscale parts. These micro or nanoscale robots would allow nanometer precision which would permit doctors to perform curative and reconstructive procedures at the cellular and molecular level. Scientists hope to develop nanomachines that could monitor and transmit information about a patient’s internal body chemistry, warning doctors of any chemical imbalances within the circulatory and lymphatic systems. They hope to develop nanomachines that could be implanted into the nervous system and monitor pulse, brain-wave activity and other neural functions. They also hope to be able to develop nanodevices that could dispense drugs and hormones, only when and where needed, in individuals with chronic imbalances and deficiencies. Scientists hope to be able to someday devise artificial antibodies, white and red blood cells and antiviral nanorobots. The development and use of these nanodevices and nanorobots can only be accomplished when the design and construction the devices and/or their components can be achieved. This has not been easily accomplished. The construction of these nanomachines and nanorobots is being actively pursued, and there have been some exciting advancements in this area of nanomedicine. However, these nanomachines are nowhere near the level of applicability of some of the delivery system nanoparticles and imaging nanoparticles that will be discussed below.
History of Nanotechnology
1959 – The concept of nanotechnology began with the physicist Richard Feyman. He explored this concept of the possibility of controlling matter at the nanoscale level. He described the idea of having the entire Encylopedia Britannica written on the head of a pin.
Early 1970’s – IBM began using electron beam lithography to create nanostructures and nanodevices as small as 40-70 nm. Electron beam lithography creates the extremely fine lines in the integrated circuits patterns seen in electronics.
Late 1970’s – Eric Drexler began formulating the concept of molecular nanotechnology.
1974 – The term “nanotechnology” was used for the first time by Norio Taniguchi. It was used in reference to the ability to engineer materials at the nanometer scale. At this time, the driving force behind “nanotechnology” was electronics. The goal of this miniaturization was the development of tools for making smaller, and therefore faster, electronic devices on silicon chips.
1986 – Eric Drexler released his book entitled, Engines of Creation: The Coming Era of Nanotechnology, which described molecular nanotechnology. Molecular nanotechnology is the concept of engineering functional mechanical systems on the molecular scale. Drexler’s book describes the idea of a nanoscale “assembler” which would be able to build a copy of itself and of other items of arbitrary complexity. Not everybody fully supports Drexler as some of his ideas are unrealistic.
The Use of Nanostructures for Drug and Gene Delivery
Many of the orally administered, and even intravenously administered, drugs used today have many downfalls and disadvantages. The problems associated with these drugs include:
– Alteration or destruction of the drug due to the acidic conditions within the stomach.
– Alteration or destruction of the drug due to the first-pass effect (presystemic elimination); drugs usually enter the circulation via the hepatic portal system, and can therefore be metabolized by the liver before they enter the systemic circulation.
– Resistance by the intestine, resulting in low absorption of the drug.
– Drugs that circulate systemically can become toxic to non-target tissues along with the target tissue.
– Many drugs are unable to cross the blood brain barrier.
– May be difficult to maintain a therapeutic concentration of the drug.
– The drug may be metabolized within the systemic circulatory system or within the interstitium.
– The drug may be rapidly excreted.
– The drug may have limited solubility within the circulatory system.
– The drug may have limited cell permeability.
– The drug may require toxic solvents or adjuvants to enhance its solubility or activity.
Many of the chemotherapy drugs used today peak at a concentration greater than the maximum tolerable level, and then the concentration quickly plummets below the effective therapeutic level. This requires high doses and frequent administration of the chemotherapy drugs, which patients tend to not be able to tolerate for extended periods of time. It is not only with chemotherapy drugs that it is difficult to achieve and maintain effective therapeutic levels, but this is also the case with many other types of drugs that are used to treat many other types of diseases and disorders.
The ideal drug delivery system should be able to accomplish the following:
– high stability of the delivery system
– high absorption of the delivery system
– high therapeutic concentration at the target site and not simply systemically
– long-term release of the drug at the target site
– low administration frequency
– protection of the drug to be delivered
Some of the recently developed nanostructures possess many of these qualities, and are therefore being further studied for their future use as drug delivery systems, and some have already been approved for use as delivery systems in humans. Many of the nanoparticles (NP’s) have been developed to encapsulate the drug. Some of these NP’s have additional applications, including gene delivery and imaging. Nanostructures are used for the exploitation of small anatomical and physiological differences that occur at the molecular level between diseased and normal tissues. These small differences are used as a way of targeting drug and gene delivery to diseased tissues of interest. The various types of NP’s and their applications in nanomedicine will be further discussed below.
Liposomes are the simplest form of NP. They are comprised of a spherical phospholipid bilayer. This bilayer is formed from non-toxic, natural phospholipids and cholesterol, similar to the phospholipids bilayers that surround our cells. Liposomes are variable in their composition and are therefore very versatile in their applications. They can be loaded with a variety of different molecules including: small drug molecules, proteins, nucleic acids, and even plasmids.
Figure 1. This diagram demonstrates the variability in liposome composition.
The variability of liposomes stems not only from the different molecules that they can encapsulate, but also what type of modifications are made to the outer surface of the liposome. The outer surface of liposomes can be modified such that the surface is more hydrophilic. This significantly increases the circulation time of liposomes in the bloodstream. These “stealth” liposomes are being used in chemotherapy treatment for cancer. These liposomes have been used as carriers for hydrophilic chemotherapy drugs such as doxorubicin, mitoxantrone and others.
Specific antibodies or binding proteins can also be conjugated to the outer surface of drug-carrying liposomes to improve the specificity of drug delivery. An antibody specific to a certain receptor that is upregulated in the tumor cells, or specific to an antigen that is specific to a certain type of cancer cell, can be used to target the drug-carrying liposome to these target tumor cells.
Figure 2. Binding of targeted liposomes to receptors on the outer surface of a cancer cell.
Liposomes can encapsulate hydrophilic drugs, confining and protecting the drug until the liposome binds to the target cell. However, in the case of lipophilic and amphipathic drugs, these drugs become tightly incorporated into the phospholipid bilayer of the liposome. This increases the solubility of the amphipathic or lipophilic drug, and at the same time protects it.
Figure 3. How amphipathic drugs become incorporated into the lipid bilayer of a liposome.
The blood brain barrier is a major obstacle in the delivery of drugs to the brain. The endothelial cells that line the capillaries in the brain form the blood brain barrier. These endothelial cells are tightly packed via tight junctions, blocking the movement of all molecules out of the capillaries into the brain, except for those that are small and lipid soluble (gases, ethanol and steroids), and those with specific transport systems (glucose and amino acids). Even if a molecule does enter an endothelial cell within the blood brain barrier, this does not guarantee its entry into the brain. The endothelial cells of the blood brain barrier can metabolize certain molecules, preventing their entry into the brain. Only those molecules and nutrients required by the brain for survival are allowed to cross the blood brain barrier. There are four mechanisms by which molecules can cross this barrier: passive diffusion, carrier-mediated transport, carrier-mediated transcytosis, and absorptive-mediated transcytosis.
Liposomes have been shown to be able to cross the blood brain barrier. A receptor-specific monoclonal antibody conjugated to a PEGylated liposome was used to target liposomes to a rat transferring receptor, which is abundant in the microvascular endothelium of the brain. These antibody conjugated liposomes resulted in greater delivery of the drugs to the brain.
PEG can be conjugated to the outer surface of most NP’s. NP’s are commonly PEGylated to increase their biocompatibility. PEGylation decreases the immunogenicity of NP’s. Macrophage, which are part of the innate immune system, do not recognize PEG molecules as foreign. PEGylation decreases the uptake of the NP by macrophages in the reticuloendothelial system (spleen and liver). This results in an increase in the circulation time of the NP and drug, gene or imaging molecule being delivered.
Figure 4. Poly(ethylene glycol) (PEG)
One of the most recently developed liposomes consists of the conjugation of a cell-penetrating peptide called the TAT peptide and PEGylation for protection. PEG2000 molecules were attached to the phosphatidylethanolamine phospholipid via a pH-sensitive hydrazone bond. PEG1000 molecules were used to attach the TAT peptide to the surface of the liposome. At a close to neutral pH, the PEG2000 molecules shield the TAT peptide due to the shorter PEG1000 spacers. These liposomes, along with many other carrier systems, accumulate spontaneously within tumor and ischemic tissues due to the enhanced permeability and retention effect. Once the lipsomes arrive to the acidic environment within tumor or ischemic tissues, pH-induced hydrolysis of the hydrazone bonds occurs. This causes the protective PEG shield to be lost and the TAT peptides to become exposed. The exposed TAT peptides allow the liposomes to penetrate even further for specific drug delivery into the tumor and ischemic cells. These liposomes have been shown to be effective for specific drug delivery in in vitro cell cultures, and also in vivo, in ischemic cardiac tissues in rat heart models and in mice models.
The advantages of using liposomes as a drug delivery system compared to administration of the drug alone include:
-longer circulation times of the drug
-decreased systemic toxicity
-increased uptake of the drug by the target tissue
-steady release of the drug (leads to a maintained therapeutic concentration)
Figure 5. A polyamidoamine (PAMAM) dendrimer
Dendrimers are comprised of polymers which form branches around a central core. They can be made from many different types of polymers including: poly(L-glutamic acid) (PGA), polyamidoamine (PAMAM), poly(ethylene glycol) (PEG), and polyethylenimine (PEI). Each dendrimer consists of a multifunctional core molecule with a dendritic wedge attached to each its functional sites. The molecular structure of dendrimers is highly uniform, there is a narrow molecular weight distribution, they can be formed to be a specific size and to have specific shape characteristics, and they have highly functionalized termini. Dendrimers are so uniform and can be made to be a specific size because they are built up one layer at a time, or one “generation” at a time.
Figure 6. The synthesis of a PAMAM dendrimer.
The mechanism involved in the synthesis of the commonly used PAMAM dendrimer involves a conjugate addition of PAMAM to the α,β unsaturated carbonyl, methyl propenoate, followed by a substitution reaction between the product and another PAMAM molecule.
The highly functionalized termini of dendrimers form a vast amount of surface area to which therapeutic agents and targeting molecules can be attached. These termini are most often composed of amine or acid groups which provide the means through which many different types of functional components can be attached. These termini can be linked to drugs, nucleic acids, or imaging molecules, such as fluorescent dyes. The termini of dendrimers can also be covalently linked to specific antibodies or peptide binding motifs for specific targeting of drug, gene, or imaging molecules. Their highly charged surfaces make them highly soluble in aqueous solutions such as those found in the blood and interstitial tissues. Their charged surfaces also allow them to bind to cell surfaces. It has been shown that dendrimers have an in vitro size-dependant cytotoxicity. The more generations of the dendritic branches, the greater the surface charge, and therefore the greater cell membrane adherence. Encapsulation of drugs inside dendrimers has been successful, however, the release rates were extremely slow. Dendrimers are most effective when the therapeutic agent is conjugated to the termini of the dendrimer.
PAMAM dendrimers are very efficient at penetrating the intestinal epithelial cells and have thus been shown to greatly enhance the bioavailability of many types of drugs. A specific example is a recently developed PAMAM dendrimer that is conjugated to the model chemotherapeutic drug doxorubicin. Transport from the mucosal side to the serosal side of the small intestine of rats was shown to be 4-7 times greater with treatment of the doxorubicin-PAMAM complex compared to administration of the free doxorubicin drug. After oral administration, the bioavailability of the doxorubicin-PAMAM complex was 200-fold greater than the free doxorubicin. This research demonstrates the promise of PAMAM dendrimers in the enhancement of the bioavailability of P-glycoprotein substrates in particular.
These NP are most commonly formed from polylactic acid (PLA) or poly(lactic-co-glycolic acid) (PLGA) polymers.
Figure 7. Polylactic acid molecule.
Figure 8. Poly(lactic-co-glycolic acid) (PLGA)
Figure 9. Scanning electron microscopy of a NP formulation.
PLA and PLGA are used in the production of NP’s because these polyester polymers are biodegradable. The degradation of these polymeric NP’s within the body causes release of the encapsulated drug or DNA, and the metabolites produced, lactic acid and glycolic acid, are easily removed through the citric acid cycle during cellular respiration. Not only are the polymeric NP’s biodegradable, but they have many other advantages including: sustained release, which results in a maintained therapeutic concentration; biocompatibility, which means that the NP’s are not immunogenic, are non-toxic and perform their intended function; and they have the ability to protect the therapeutic agent that they encapsulate from degradation. Also, in terms of biocompatibility, not only are the NP’s themselves non-toxic, they eliminate the requirement for the toxic solvents and adjuvants that are sometimes required to increase the solubility and activity of drugs.
Many of the molecules that are brought into our cells go through receptor-mediated endocytosis. It is the binding of the ligand molecule to the receptor that signals the initiation of endocytosis of the ligand molecule. After binding of the ligand to the receptor, these ligand-receptor complexes group together in clathrin-coated pits. These clathrin-coated pits continue to invaginate until they finally pinch off forming a clathrin-coated vesicle within the cell’s cytoplasm. This clathrin coat must be removed so that the vesicle can fuse with other membranes. Once they are uncoated, these individual vesicles can fuse forming early endosomes. These early endosomes can then fuse with lysosomes forming an endolysosome. Lysosomes contain digestive enzymes such as acid hydrolases. A lysosome fuses to an early endosome so that degradation and recycling of the receptor can occur.
It has been shown that NP’s rapidly escape from these endolysosomal compartments following receptor-mediated endocytosis, thus protecting the NP’s and encapsulated therapeutic agents from the degradative environment within the endolysosome.
Figure 10. Receptor-mediated endocytosis
Functionalization of the biodegradable NP’s involves the modification of the outer surface of the NP’s. These modifications involve the conjugation of antibodies, peptides, specific aptamers, or other biological molecules to the surface of the NP’s. These molecules will interact with specific tumor or disease-associated antigens and receptors, allowing targeted delivery of the NP’s to the tissue of interest.
Aptamer conjugated-NP’s have shown a high level of specificity for drug delivery to certain prostate cancer cells. A PEGylated PLGA copolymer was used to encapsulate the chemotherapeutic drug docetazel (Dtxl). This NP was then conjugated to the A10 2’-fluoropyrimidine RNA aptamer. The aptamer was conjugated to the termini of the PEG molecules via carbodiimide coupling chemistry between the carboxylic acid moiety at the terminus of the PEG molecule and the amine-functionalized aptamer. The carbodiimide activates the carboxylic acid towards nucleophilic attack by the amine.
Figure 11. A carbodiimide coupling reaction.
The A10 2’-fluoropyrimidine RNA aptamer binds specifically to the extracellular domain of the prostate-specific membrane antigen (PSMA). PSMA is a well-characterized antigen that is expressed in prostate cancer cells. The binding of the NP-conjugated aptamer to PSMA allows specific delivery of the chemotherapeutic drug to the prostate cancer cells. In vivo studies have shown that these drug-encapsulated NP-aptamer bioconjugates are very effective at inhibiting the growth of the prostate cancer cells and show decreased systemic toxicity compared to administration of the drug alone.
Just as specific antigens and upregulated membrane proteins can be exploited to target NP’s to cancer cells, there is other pathophysiology of tumor tissues that can be exploited for targeted drug and gene delivery. Tumor tissues commonly show: increased angiogenesis, hypervasculature, defective vascular architecture, impaired lymphatic drainage, and an acidic environment due to lactic acid build-up. All of these pathophysiological characteristics can be used to help target NP’s to tumor tissues. Poly(ethylene oxide) (PEO)-modified poly(β-aminoester) (PbAE) was used to form pH-sensitive NP’s. These NP’s were used to encapsulate the chemotherapeutic drug paclitaxel. Tumor tissues are commonly hypoxic. This leads to an accumulation of lactic acid due to anaerobic cellular respiration. The low pH caused by the lactic acid in the tumor tissue allows for selective release of the paclitaxel from the pH-sensitive NP’s into tumor cells. In normal chemotherapy, the aqueous paclitaxel solution is administered intravenously and is fairly toxic. Therefore, these NP’s not only improve the therapeutic efficiency by selectively releasing the drug in the target tissue, but they also decrease the systemic toxicity of the drug. There is an enhanced permeability and retention effect within tumor tissue as a result of the defective vasculature and impaired lymphatic drainage. The increased permeability of the blood vessels within tumor tissues leads to increased extravasation (leakage of fluid, molecules, and cells from the blood vessels). This is due to the large pores observed in the blood vessels within tumor tissue. These pore sizes have been observed to be between 380 and 780 nm, and NP’s are commonly between 100 and 200 nm. The increased retention is caused by the decreased lymphatic drainage. Both of these characteristics allow the passage of the NP’s into the extravascular spaces and their accumulation within the tumor tissue. It has been experimentally determined that NP’s accumulate within tumor tissues to a level 6-fold higher than within normal tissues. However, the problem with this approach is that not all blood vessels within tumors show these same leakage and retention properties.
Researchers are now looking at dual-ligand approaches for drug and gene delivery. The conjugation of two different types of ligands to the outer surface of the NP’s increases the selectivity of the NP for the target tissue. Many of the ligands that are commonly used bind not only to the target cells, but also to many other healthy cells in the body. By increasing the number of ligands conjugated to the NP’s, the specificity of the NP’s can be increased, thus decreasing the binding of the NP’s to non-target cells.
NP’s are commonly formed via a nanoprecipitation reaction. This reaction involves only one step with the NP’s forming instantaneously. The reaction requires two solvents that are miscible; one in which the drug and the polymer are soluble (solvent), and the other in which they are not (non-solvent). Nanoprecipitation occurs due to the rapid desolvation of the polymer as the solvent diffuses into the non-solvent. When the solvent is added to the non-solvent (also called the dispersing medium), the polymer precipitates leading to immediate NP formation and drug entrapment.
Nanoprecipitation occurs due to the Marangoni effect. The Marangoni effect is caused by interfacial turbulence between two solvents which is the result of the complex and cumulated phenomena of flow, diffusion and surface tension variations. Nanoprecipitation has many advantages. There is no extended stirring, sonication or high temperatures required, and there are no oily-aqueous interfaces formed, which can all cause protein damage. There are also no toxic solvents or surfactants required.
The most common solvent and non-solvent used are ethanol and water respectively. This solvent, non-solvent combination does show 100% entrapment efficiencies, however, they are only suitable for hydrophobic drugs. Using PLA and PLGA as polymers, it has been shown that dimethyl sulfoxide (DMSO) is the most favorable solvent for encapsulation of protein drugs, and polar aprotic, ketones and ester solvents are most effective for encapsulation of hydrophilic drugs. The non-solvent is chosen based on its polarity to favor final drug loading. Finding the right solvent, non-solvent pair is a matter of screening different combinations. However, there are a couple of concepts that determine favorable solvent, non-solvent combinations. The greater the rate of diffusion of the solvent into the non-solvent, the smaller the size of the NP’s, the greater the yield of NP’s, and the greater the percentage of encapsulation.
The rate of diffusion of the solvent into the non-solvent is affected by the difference in dielectric constant between the two solvents. The dielectric constant refers to the ability of a material to store a charge from an applied electromagnetic field and then transmit that energy. The more hydrophilic a compound, the higher its dielectric constant. The smaller the difference between the two dielectric constants of the solvent and the non-solvent, the greater the rate of diffusion of the solvent into the non-solvent because of the higher affinity of the solvent for the non-solvent.
It is also important to ensure that the concentration of the polymer within the solvent is not too high. A high polymer concentration can inhibit NP formation. However, the ratio of volumes of the solvent and the non-solvent does not affect NP formation and can therefore be altered to optimize the yield of the NP’s. When comparing solvents of homologous compounds, the greater the number of carbon atoms in the backbone of the compound, the more destabilized the NP.
NP recovery after the nanoprecipitation reaction can sometimes be difficult. If the non-solvent is organic, high-speed centrifugation leads to caking of the NP, while low-speed centrifugation leads to low yields. The non-solvent can be replaced with water, however this tends to lead to coalescence of the NP, altering their morphology.
Nanocells are comprised of a nuclear PLGA NP, surrounded by a PEGylated phospholipid envelope. These nanocells are preferentially taken up into tumor tissues just as the simple biodegradable polymeric NP’s were, again due to the enhanced permeability and retention effect. However, because of the double layer of the nanocells, one chemotherapeutic drug can be encapsulated within the NP, while a different chemotherapeutic drug can be trapped within the phosolipid envelope. This allows the temporal release of the two drugs. This temporal release has been investigated in murine tumor models with doxorubicin encapsulated within the NP and combretastatin-A4 trapped within the lipid envelope. The degradation of the envelope releases the first drug leading to vascular shutdown. The degradation of the NP within the tumor leads to the release of the second drug which causes vascular collapse. These experiments showed that the combination of the two drugs was more effective in suppressing tumor growth than a single drug delivery therapy.
Figure 12. A nanocell: composed of a NP core surrounded by a PEGylated phospholipid bilayer
This is a newly patented technology that uses magnetic nanoparticles and targeted magnetic fields to control cell activity. The core of the nanoparticle is comprised of a magnetic material. The magnetic core is surrounded by a protective shell which minimizes corrosion of the core. An organic molecule acts as the linker to conjugate the biomolecule (a specific protein binding motif) to the protective shell.
Figure 13. General components of a magnecell
The protein binding motifs conjugated to the outside of the magnecells can bind to ion channel receptors or mechanosensitive channel receptors. When a magnecell binds to a specific receptor via the protein binding motif and a magnetic field is applied, the magnetic magnecells are displaced causing the channels to which they are bound to open. This leads to activation of various cell signaling pathways, leading to alterations in gene expression and other cellular responses.
Figure 14. This diagram illustrates the binding of the magnecell to the cell membrane. When the magnetic field is applied, the displacement of the magnecell causes a slight deformation in the membrane and opening of the mechanosensitive channel.
Figure 15. This diagram illustrates the binding of the magnecell to a specific ion channel receptor. When the magnetic field is applied, the displacement of the magnecell causes the ion channel to open.
This technology allows the opening of specific channels and the subsequent activation of specific signal transduction pathways without the use of drugs or other biochemical stimuli. Magnecells have already been shown to aid in the production of cartilage by controlling the activation and differentiation of stem cells into chondrocytes; this has been shown both in vitro and in vivo. Magnecells have also been shown to control the activation of mechanosensitive potassium channels and calcium channels. Magnecells show the potential to be able to control many other types of channels and will potentially have many applications in the control and targeted therapy of many diseases in which faulty ion channel function is to blame; for example Cystic Fibrosis.
The initial technique used for gene delivery was via viral vectors. However, there are many disadvantages to the use of viral vectors. The patient can develop a large immune response against the viral proteins. The viral vectors insert into the genome randomly, and therefore have the potential to become oncogenic. If the viral vector inserts into a proto-oncogene, producing an oncogene, the patient can potentially develop cancer. There is also the potential of the viral vector reverting back, via random mutation, to its original virulent form.
Liposomes, dendrimers and polymeric NP’s can all be used for gene delivery. The use of these various types of NP’s for gene delivery eliminates many of the problems caused by the use of viral vectors. NP’s have little to no immunogenicity. Unlike viral vectors, gene delivery via NP’s does not require the introduction of extra genes, and therefore no potentially immunogenic or toxic proteins are introduced into the patient. Viral vectors are limited in the size of the gene that they can carry. However, there is no limit to the size of the gene that can be delivered and transferred via NP’s. The DNA, RNA, or siRNA can be encapsulated within NP’s or the negatively charged backbone can interact electrostatically with positively charged surfaces of NP’s.
NP based assays have been shown to be able to detect proteins in the attomolar range (10exp-18). Enzyme-linked ImmunoSorbent Assay (ELISA), a commonly used biochemical technique for the detection of specific antibodies or antigens in a sample, can only detect in the picomolar range (10exp-12). To put this into perspective, if a molecule had a concentration of 100 attomolar, there would be less than 1 of these molecules per cell. This NP based biosensing will allow for early detection of cancer and other diseases, early treatment and will hopefully be able to save many lives.
Quantum dots, also known as nanocrystals, are commonly composed of a CdSe core, which is capped with several monolayers of ZnS. ZnS has a much larger bandgap than CdSe and eliminates inherent surface defects of the CdSe core, thus increasing quantum yield. The ZnS shell allows a brighter emission from CdSe by reducing non-radiative recombination. Non-radiative recombination occurs when the excited electron recombines with its hole but does not produce electromagnetic radiation. Many methods have been developed for conjugating biologically important molecules to the ZnS surface. Because of the CdSe core composition, quantum dots are semi-conductors. A quantum dot consists of 10-50 atoms and has a diameter of 2-10nm, making it a nanoparticle. Quantum dots have many desirable characteristics including: sharp and symmetrical emission spectra, high quantum yield, broad absorption spectra, good chemical and photostability and size-dependent wavelength emission tunability. Many of these features give quantum dots an edge over traditional dyes. In addition, traditional dyes, organic imaging compounds and radioactive tags tend to have a much shorter lifespan than quantum dots; quantum dots have been found to remain in animal systems for months.
The usefulness of quantum dots lies in their ability to emit light at very specific wavelengths. The color of light emitted by a quantum dot is dependent on its size. The larger the quantum dot, the redder the emission. Quantum dots of a 2nm diameter emit bright green light, while one with a 5nm diameter yields red light. The color of the emission is dependent on the size of the quantum dot because the size of the quantum dot determines the size of the band gap. A larger band gap means a longer fall for an electron and therefore a greater amount of energy released. It is easy to adjust the band gaps of quantum dots as you simply add more atoms to the core to alter the band gap boundaries. The band gaps of traditional semi-conductors cannot be easily adjusted in this manner.
Figure 16. CdSe core with ZnS shell quantum dots with diameters from 1.9nm (blue emission) to 5.2nm (red emission). This graph highlights the wide adsorption spectrum and narrow emission spectrum. This data is from quantum dots constructed by EvidentTech.
The main biomedical application of quantum dots is imaging. Imaging works by shining light, which can be white or near-infared, and this light gets absorbed by the quantum dots causing an electron to be excited from the valence band to the conduction band. When this electron falls back down into the hole left in the valence band, light is emitted at a specific wavelength. Kim et al. have shown that quantum dots can be used to do sentinel lymph node mapping at a tissue depth of 1cm. This application could lead to improvement of the sensitivity of lymphatic resectioning surgery. This surgery is used to reduce metastasis in some cancers. One of the principle features of quantum dots that makes them so useful in imaging applications is their ability to be tuned to emit different color spectra. Lidke et al. have used quantum dots to begin to map signal transduction pathways in cells. This can be done through observing co-localization of different groups of quantum dots that have been targeted to different proteins and that are emitting different wavelengths of light. When used as molecular imaging probes, quantum dots provide very high resolution viewing of cellular components and activities. Imaging of tumors has been carried out using quantum dots which have been immunolabeled with HER2 cancer markers. Because of the long lifespan of quantum dots they can be used to track tumor metastasis after they have been targeted to a tumor.
Figure 17. Quantum dots that have been labeled with an antibody for over-expressed Her2 recpetors on cancer cells. These cancer cells have a blue nuclear staining.
Another medical application of quantum dots is the use of labeled quantum dots to carry silencing RNA (siRNA). Quantum dots can be conjugated to human epidermal growth factor. The human epidermal growth factor receptor is highly overexpressed in certain cancer cells. A quantum dot labeled with this growth factor will undergo receptor mediated endocytosis. If this quantum dot also has siRNA conjugated to it, then the siRNA will be delivered to cancer cells as well. This gives rise to dual functionality of the quantum dot. Now the tumor cells, which have taken up the quantum dots, will be affected by the siRNA, and can be imaged. This provides a way to see where the siRNA has been carried within the body.
Quantum dots show promise in the field of nanomedicine, but one final issue must be addressed. Studies have shown that quantum dots show significant cytotoxic effects on the body. When the CdSe core is oxidized it releases Cd2+ ions which are highly toxic to the liver. This issue must be resolved before quantum dots can take over as the new answer to medical imaging. Quantum dots are also not very cheap. On the Evident Technologies website they are listed as $449 for 50mg. You can save money by buying in bulk; 200mg will only set you back $749.
Figure 18. A spectrum of quantum dot emissions, with quantum dot size increasing from left to right (bandgap getting smaller left to right).
Nanoshells are most commonly composed of a silica core and a gold shell. Gold is a good choice for the shell material as it is generally very biocompatible and inert. Nanoshells can be made even further inert by the addition of stealth molecules such as PEG. The silica core is 120nm in diameter with a 10nm thick gold shell. The size of the core and thickness of the shell can be altered, causing the nanoshells to absorb and scatter different wavelengths of light throughout the visible and near-infared (NIR) spectrum. This adjustability allows nanoshells to be used in therapeutic and imaging applications simultaneously. The most useful nanoshell dimensions have proved to be a 120nm core with a 10nm thick shell. This size nanoshell will absorb light with an 800nm wavelength, which is in the NIR range. Nanoshells offer improved photstability compared to traditional NIR dyes due to their rigid metallic structure.
Figure 19. Nanoshells with a silica core and gold shell shown have their optical resonance peak tuned to 975nm, which is in the NIR range.
The medical applications of these nanoparticles are thermal ablation therapy and imaging. Nanoshells can be conjugated to antibodies against proteins that are known to be over-expressed on certain types of cancer cells. In one specific case this was done by attaching anti-HER2 antibodies to a polyethylene glycol (PEG) linker, and then the PEG linker/antibody complex was attached to the surface of the nanoshell. The antibody is attached to the PEG linker through an amidohydroxysuccinimide group. The antibody/PEG linker complex is attached to the surface of the nanoshell through a sulfur group at the opposite end of the complex. Cancer cells over-expressing HER2 will bind to this complex. These cells can then be located via nanoshell imaging and killed via thermal ablation. This has been carried out in vitro.
Imaging is a less prominent feature of nanoshells compared to the their potential for thermal ablation therapy. The imaging potential of nanoshells is not as great as that of quantum dots but is still useful as it provides a dual functionality of the nanoshells; the ability to image and destroy cells. Thermal ablation therapy is where nanoshells show the most promise. Nanoshells can absorb light in the NIR range (800nm). When this wavelength of light is shined on tissues it will penetrate through the tissue without harming it, but when the light reaches the nanoshells it causes them to heat up. The nanoshells can create sufficient heat to kill the cell in which they are contained. This is the process of thermal ablation therapy.
Nanoshells have also been used to take advantage of the enhanced permeability and retention effect displayed by tumors. An in vivo study was carried out by O’Neal et al. in which nanoshells were targeted to tumors in mice solely through this enhanced permeability and retention effect. This method of targeting was sufficient to provide therapeutic effects (i.e. kill the tumors) while not harming the surrounding tissues. This provides a solution to previous thermal ablation treatments, which employed a laser, but could not discriminate between normal cells and tumor cells.
Nanoshells have not proved to be toxic up to this point. Their non-toxicity is likely due to the use of gold as a shell,which as mentioned earlier, is biologically inert.
Carbon nanotubes were discovered in the past 50 or so years. However, it is not clear when they were first discovered. A Russian journal published evidence of carbon nanotubes in the late 1950’s but this discovery went mostly unnoticed. In the 1970’s a few more researchers were successful in creating nanotubes, and by the late 1980’s they were known as a distinct molecular form of carbon atoms. Nanotubes exhibit many different characteristics, they can be single walled or multi-walled, and conducting or semi-conducting. Nanotubes are made of carbon which take on same arrangement as in graphite, only rolled into sheets. They have a diameter on the order of nanometers but are generally very long; 50 microns would be a typical length. Researchers at the University of Cincinnati have developed methods for making carbon nanotube arrays of up to 18mm in length. Nanotubes are 100x as strong as steel while being only 1/6 the weight. The heterogeneity and aggregation of tubes creates difficulties in the use of nanotubes as sensors as a homogenous set of nanotubes should be used.
Figure 20. A single walled carbon nanotube seen from the side.
Figure 21. A single walled carbon nanotube seen from the front.
Carbon nanotubes absorb light in the NIR range. This allows them to be used in thermal ablation therapy in a manner very similar to that of nanoshells. One study used carbon nanotubes in thermal ablation therapy by putting folic acid on the outside of the nanotube. Folic acid is a vitamin and cancer cells often over-express the folate receptor and will therefore take up these labeled nanotubes through receptor mediated endocytosis. Single walled carbon nanotubes have been shown to generate large amounts of heat at relatively low concentrations when exposed to NIR light. One study showed that by shining NIR radiation on a solution containing nanotubes at a concentration of 25mg/L (suspended in water), the solution could quite easily be brought to boiling. It is easy to see how these nanotubes could be used in the same way as nanoshells to carry out thermal ablation therapy. And once again, any cells that have not taken up the carbon nanotubes will be unaffected by the NIR radiation. This allows selective destruction of cancer cells.
Another interesting application of nanotubes is the delivery of DNA to cells. Nanotubes with a 1-2nm diameter have been shown to be able to carry pieces of DNA 15 bases in length absorbed to their surfaces. These nanotubes have been shown to be taken up by cells at a rate which increases with increasing temperature, indicating an energy dependent uptake mechanism. Once these nanotubes are taken up by cells, short pulses of NIR radiation will cause the release of the DNA from the surfae of the nanotube. This DNA will then be translocated to the nucleus of the cell. There was no apparent cytotoxic effects on the cells which were receiving and accumulating these nanotubes in their cytoplasm.
Another application of nanotubes is their use to deliver siRNA to cells. The nanotubes can be targeted to cancer cells and down-regulate targeted genes. Nanotubes have been designed with –CONH-(CH2)6-NH3+ functional groups on their surface. This positive charge mediates the conjugation of siRNA. Previously siRNA delivery was largely unsuccessful as degradation would occur before delivery of the siRNA to the targeted cells. Nanotubes seem to be able to complete this delivery with the siRNA still intact.
Nanotubes are semiconductors and this property has been exploited in biological applications, specifically in detetecting mercuric ion levels in various mammalian fluids and tissues. Nanotubes have a very high photostability and a band gap that is easily altered by the surrounding environment. In a study by Heller et al., double stranded DNA 30 basepairs in length was absorbed to the outside of single walled carbon nanotubes. This DNA was put into solution with carbon nanotubes and spontaneously absorbed to the surface of the nanotubes in a double stranded, right handed, native B form. The DNA was attached to the nanotube through non-covalent binding. When mercuric cations were present, they absorbed to the DNA shielding the negative charges from one another. This causes the DNA to undergo a conformational change and take on the Z form. This changes the electronic environment around the nanotube and decreases its emission energy. Depending on the number of ions that bind the DNA the emission energy can be changed by up to 15meV; this allows varying concentrations of ions to be detected. The binding process is reversible which makes these nanotubes reusable. The emission energy will return to its original value upon removal of the ions, and the DNA will flip back to its B form. This sensing has been carried out in whole mammalian blood and living mammalian cells.
Figure 22. DNA wrapped around a single walled carbon nanotube changing from B form to Z form as cations bind.
Another issue that must be addressed with nanotubes is toxicity. Studies have shown that single walled carbon nanotubes linked with larger molecules tend to show significant cellular toxicity. These are nanotubes conjugated with things such as streptavidin-biotin, a large complex. This type of nanotube tends to exhibit extensive cell death. There is conflicting information on whether or not carbon nanotubes are toxic on their own. Cui et al. have stated that nanotubes, without any other molecules conjugated to them can inhibit cell proliferation and down-regulate cell cycle genes, inducing apoptosis. Meanwhile Kam et al. have stated that single walled carbon nanotubes are not toxic unless they are conjugated with larger molecules. More research must be carried out in this area to determine whether or not carbon nanotubes are safe for medicinal use.
Supraparamagnetic particles are Fe3O4 particles (known as iron oxide or magnetite) less than 10nm in diameter. These particles have long been used in magnetic resonance imaging (MRI). These nanoparticles are often formed by alkaline coprecipitation using Fe2+ or Fe3+ salts in water with a hydrophilic polymer. This hydrophilic polymer can be dextran or polyethylene glycol (PEG). The end result is a crystal with an iron core, 4-5nm in diameter, and a coating consisting of the polymer. As mentioned above, applying a PEG coating makes molecules, including nanoparticles, less immunogenic and more stealthy in the body. This increases their circulation time and biocompatibility. This crystal has a large magnetic moment when brought into a magnetic field and they can induce a loss of signal intensity from protons. The surface of these particles can be modified through the addition of oligonucleotides, proteins, or antibodies. These particles can once again be used to treat cancer through thermal ablation therapy. They can generate sufficient heat to kill cells through the application of a magnetic field. Brownian relaxation of the particles occurs during the application of an alternating magnetic field and this generates the heat through particle rotation in the field.
Johannsen et al. have used superparamagnetic nanoparticles to treat prostate cancer. The treatment involved injection of the nanoparticles into the tumor mass transperineally. The particles were in solution at a concentration of 120mg/ml and 12.5ml of this solution was injected over 24 injection sites. An alternating magnetic field was applied and intra-tumoral temperatures of up to 48.5 degrees Celsius were reached. For thermoablative therapy to be effective, temperatures of 42 degrees Celsius must be reached, which they were in this study. Another benefit of treating tumors with these nanoparticles is the very low clearance rates of the tumor mass. This means that the nanoparticles injected into the tumors will stay there for long periods of time. In the study mentioned, the patient had computed tomography (CT) scans done six weeks after the nanoparticle injection and the deposits of nanoparticles were still clearly visible. This is beneficial as it provides the oppourtunity to carry out multiple treatments after only one injection treatment. This avoids multiple uncomfortable transperineal injection experiences for the patient. This study demonstrates that hyperthermic treatment of prostate cancer with nanoparticles is feasible. Toxicity studies are to be performed following this study, but so far these nanoparticles do not appear to be toxic.
Figure 23. A patient undergoing thermal ablation therapy using superparamagnetic nanoparticles for the treatment of prostate cancer.
Bilati, U., Allemann, E., and Doelker, E. Development of a nanoprecipitation method intended for the entrapment of hydrophilic drugs into nanoparticles. European Journal of Pharmaceutical Sciences. 2005; 24(1): 67-75
Johannsen, M., Gneveckow, U., Eckelt, L., Feussner, A., Waldofner, N., Scholz, R., Deger, S., Wust, P., Loening, S.A., Jordan, A. Int. J. Hyperthermia. 2005; 21(7): 637-647
Jin, S. and Ye, K. Nanoparticle-mediated Drug Delivery and Gene Therapy. Biotechnology Programs. 2007; 23: 32-41
Kale, A.A. and Torchilin, V.P. “Smart” Drug Carriers: PEGylated TATp-Modified pH-Sensitive Liposomes. Journal of Liposomes Research. 2007; 17(3): 197-203
Ke, W., Zhao, Y., Huang, R.,, Jiang, C., and Pei, Y. Enhanced oral bioavailability of doxorubicin in a dendrimer drug delivery system. Journal of Pharmaceutical Sciences. 2007. (Epub ahead of print)
Kelly, Y., Kim, M.A. Nanotechnology platforms and physiological challenges for cancer therapeutics. Nanomedicine: Nanotechnology, Biology, and Medicine. 2007; 3: 103-110
Kim, S., Lim, Y.T., Soltesz, E.G., De Grand, A.M., Lee, J., Nakayama, A., Parker, J.A., Loo, C., Lowery, A., Halas, N., West, J., Drezek, R. Immunotargeted nanoshells for integrated cancer imaging and therapy. Nano Lett. 2005; 5(4): 709-11
Mihaljevic, T., Laurence, R.G., Dot, D.M., Cohn, L.H., Bawendi, M.G., Frangiono, J.V. Nat Biotechnol. 2004; 22: 93-97
Moghimi, S., Hunter, A., and Murray, J. Nanomedicine: current status and future prospects. The FASEB Journal. 2005; 19: 311-330
O’Neal, D.P., Hirsch, L.R., Halas, N.J., Payne, J.D., West, J.L. Photo-thermal tumor ablation in mice using near-infared absorbing nanoparticles. Cancer Lett. 2004; 209(2): 171-6
Portney, N., and Mihrimah, O. Nano-oncology: drug delivery, imaging, and sensing. Anal Bioanal Chem. 2006; 384; 620-630
Sahoo, S.K., Pareveen, and Panda, J.J. The present and future of nanotechnology in human health care. Nanomedicine: Nanotechnology, Biology, and Medicine. 2007; 3: 20-31
Wust, P., Hildebrandt, B., Sreenivasa, G., Rau, B., Gellermann, J., Riess, H. et al. Hyperthermia in combined treatment of cancer. Lancet Oncol. 2002; 3(8): 487-97
http://www.evidenttech.com/ (this is an excellent website for explaining the mechanism by which quantum dots work
and explaining some applications of quantum dots)
9,509 total views, 45 views today