Developments in Bone Regenerative Medicine Using Stem Cell Treatment

Catalyst Volume 9

Abstract

There is an acute need for alternatives to modern bone regeneration techniques, which have in vivo morbidity and high cost. Dental pulp stem cells (DPSCs) constitute an immunocompatible and easily accessible cell source that is capable of osteogenic differentiation. In this study, we engineered economical hard-soft intercalated substrates using various thicknesses of graphene/polybutadiene composites and polystyrene/polybutadiene blends. We investigated the ability of these scaffolds to increase proliferation and induce osteogenic differentiation in DPSCs without chemical inducers such as dexamethasone, which may accelerate cancer metastasis.

For each concentration, samples were prepared with dexamethasone as a positive control. Proliferation studies demonstrated the scaffolds’ effects on DPSC clonogenic potential: doubling times were shown to be statistically lower than controls for all substrates. Confocal microscopy and scanning electron microscopy/energy dispersive X-ray spectroscopy indicated widespread osteogenic differentiation of DPSCs cultured on graphene/polybutadiene substrates without dexamethasone. Further investigation of the interaction between hard-soft intercalated substrates and cells can yield promising results for regenerative therapy.

Introduction

Current mainstream bone regeneration techniques, such as autologous bone grafts, have many limitations, including donor site morbidity, graft resorption, and high cost.^1,2 An estimated 1.5 million individuals suffer from bone-disease related fractures each year, and about 54 million individuals in the United States have osteoporosis and low bone mass, placing them at increased risk for fracture.^2,3,4,5 Bone tissue scaffold implants have been explored in the past decade as an alternative option for bone regeneration treatments. In order to successfully regenerate bone tissue, scaffolds typically require the use of biochemical growth factors that are associated with side effects, such as the acceleration of cancer metastasis.^6,7 In addition, administering these factors in vivo is a challenge.⁶ The purpose of this project was to engineer and characterize a scaffold that would overcome these obstacles and induce osteogenesis by controlling the mechanical environment of the implanted cells.

First isolated in 2001 from the dental pulp chamber, dental pulp stem cells (DPSCs) are multipotent ecto-mesenchymal stem cells.^8,9 Previous studies have shown that these cells are capable of osteogenic, odontogenic, chondrogenic, and adipogenic differentiation.^10,11,12 Due to their highly proliferative nature and various osteogenic markers, DPSCs provide a promising source of stem cells for bone regeneration.¹¹

An ideal scaffold should be able to assist cellular attachment, proliferation and differentiation.¹³ While several types of substrates suitable for these purposes have been identified, such as polydimethylsiloxane¹⁴ and polymethyl methacrylate¹⁵, almost all of them require multiple administrations of growth factors to promote osteogenic differentiation.⁶ In recent years, the mechanical cues of the extracellular matrix (ECM) have been shown to play a key role in cell differentiation, and are a promising alternative to chemical inducers.^16,17

Recent studies demonstrate that hydrophobic materials show higher protein adsorption and cellular activity when compared to hydrophilic surfaces; therefore, we employed hydrophobic materials in our experimental scaffold.^18,19,20,21 Polybutadiene (PB) is a hydrophobic, biocompatible elastomer with low rigidity. Altering the thickness of PB films can vary the mechanical cues to cells, inducing the desired differentiation. DPSCs placed onto spin-casted PB films of different thicknesses have been observed to biomineralize calcium phosphate, supporting the idea that mechanical stimuli can initiate differentiation.^6,16,17 Atactic polystyrene (PS) is a rigid, inexpensive hydrophobic polymer.²² As PB is flexible and PS is hard, a polymer blend of PS-PB creates a rigid yet elastic surface that could mimic the mechanical properties of the ECM.

Recently, certain carbon compounds have been recognized as biomimetic.²³ The remarkable rigidity and elasticity of graphene, a one-atom thick nanomaterial, make it a compelling biocompatible scaffold material candidate.²⁴ Studies have also shown that using a thin sheet of graphene as a substrate enhances the growth and osteogenic differentiation of cells.²³

We hypothesized that DPSCs plated on hard-soft intercalated substrates—specifically, graphene-polybutadiene (G-PB) substrates and polystyrene-polybutadiene (PS-PB) substrates of varying thicknesses—would mimic the elasticity and rigidity of the bone ECM and thus induce osteogenesis without the use of chemical inducers, such as dexamethasone (DEX).

Materials and Methods

G-PB and PS-PB solutions were prepared through dissolution of varying amounts of graphene and PS in PB-toluene solutions of varying concentrations. Graphene was added to a thin PB solution (3 mg PB/mL toluene) to create a 1:1 G-PB ratio by mass. Graphene was added to a thick PB solution (20 mg PB/mL toluene) to create 1:1 and 1:5 G-PB ratios by mass. PS was added to a thick PB solution to create 1:1, 1:2, and 1:4 PS-PB blend ratios by mass. Spincasting was used to apply G-PB and PS-PB onto silicon wafers as layers of varying thicknesses (thin PB: 20.5nm, thick PB: 202.0nm).²⁵ Subsequently, DPSCs were plated onto the coated wafers either with or without dexamethasone (DEX). Following a culture period of eight days, the cells were counted with a hemacytometer to determine proliferation, and then stained with xylenol orange for qualitative analysis of calcification. Cell morphology and calcification of stained cells were determined through confocal microscopy and phase contrast fluorescent microscopy. Cell modulus and scaffold surface character were determined using atomic force microscopy. Finally, cell biomineralization was analyzed using scanning electron microscopy.

Results

Cell Proliferation and Morphology

To ensure the biocompatibility of graphene, cell proliferation studies were conducted on all G-PB substrates. Results showed that G-PB did not inhibit DPSC proliferation. The doubling time was lowest for 1:1 G-thick PB, while doubling time was observed to be greatest for 1:1 G-thin PB. Multiple two-sample t-tests showed that the graphene substrates had significantly lower doubling times than standard plastic monolayer (p < 0.001).

After days 3, 5, and 8, cell morphology of the DPSCs cultured on the G-PB and PS-PB films was analyzed using phase contrast fluorescent microscopy. Images showed normal cell morphology and growth, based on comparison with the control thin PB and thick PB samples. After day 16 and 21 of cell incubation, morphology of the DPSCs cultured on all substrates was analyzed using confocal microscopy. There was no distinctive difference among the morphology of the DPSC colonies. DPSCs appeared to be fibroblast-like and were confluent in culture by day 15 of incubation.

Modulus Studies

In order to establish a relation between rigidity of the cells and rigidity of the substrate the cells were growing on, modulus measurements were taken using atomic force microscopy. Modulus results are included in Figure 3.

Differentiation Studies

After day 16 of incubation, calcification of DPSCs on all substrates was analyzed using confocal microscopy. Imaging indicated preliminary calcification on all substrates with and without DEX. Qualitatively, the DEX samples exhibited much higher levels of calcification than their non-DEX counterparts, as evident in thin PB, 1:1 G-thin PB, and 1:4 PS-PB samples.

After days 16 and 21 of incubation, biomineralization by the DPSCs cultured on the substrates was analyzed by scanning electron microscopy (SEM) and energy dispersive X-ray spectroscopy (EDX). The presence of white, granular deposits in SEM images indicates the formation of hydroxyapatite, which signifies differentiation. This differentiation was confirmed by the presence of calcium and phosphate peaks in EDX analysis. Other crystals (not biomineralized) were determined to be calcium carbonates by EDX analysis and were not indicative of DPSC differentiation.

On day 16, only thin PB induced with DEX was shown to have biomineralized with sporadic crystal deposits. By day 21, all samples were shown to have biomineralized to some degree, except for 1:1 PS-PB (non-DEX) and 1:4 PS-PB (non-DEX). Heavy biomineralization in crystal and dotted structures was apparent in DPSCs cultured on 1:1 G-thin PB (both with and without DEX). Furthermore, samples containing graphene appeared to have greater amounts of hydroxyapatite than the control groups. All PS-PB substrates biomineralized in the presence of DEX, while only 1:2 PS-PB was shown to biomineralize without DEX. While results indicate that while PS-PB copolymers generally require DEX for biomineralization and differentiation, this is not the case for G-PB substrates. Biomineralization occurred on DPSCs cultured on G-PB substrates without DEX, demonstrating the differentiating ability of the G-PB mechanical environment and its interactions with DPSCs.

Discussion

This study investigated the effect of hard-soft intercalated scaffolds on the proliferation and differentiation of DPSCs in vitro. As cells have been shown to respond to substrate mechanical cues, we monitored the effect of ECM-mimicking hard-soft intercalated substrates on the behavior of DPSCs. We chose graphene and polystyrene as the hard components, and used polybutadiene as a soft matrix.

By using AFM for characterization of the G-PB composite and PS-PB blend substrates, we demonstrated that all surfaces had proper phase separation and uniform dispersion. This ensured that DPSCs would be exposed to both the hard peaks and soft surfaces during culture, allowing us to draw valid conclusions regarding the effect of substrate mechanics.

Modulus studies on substrates indicated that 1:1 G-thin PB was the most rigid substrate and control thin PB was the second most rigid. In general, G-PB substrates were 8-20 times stiffer than PS-PB substrates. The high relative modulus of graphene-based substrates can be attributed to the stiffness of graphene itself. The cell modulus appeared highly correlated to the substrate modulus, both indicating that the greater the stiffness of the substrate, the greater the stiffness of DPSCs cultured on it and supporting the finding that substrate stiffness affects the cell ECM.^27,28 For example, 1:1 G-thin PB had the highest surface modulus, and DPSCs grown on 1:1 G-thin PB had the highest cell modulus. Conversely, DPSCs grown on 1:4 PS-PB had the lowest cell modulus, and 1:4 PS-PB had one of the lowest relative surface moduli. Another notable trend involves DEX; cells cultured with DEX had greater moduli than cells cultured without, suggesting a possible mechanism used by DEX to enhance stiffness and thereby osteogenesis of DPSCs.

Cell morphology studies indicated normal growth and normal cell shape on all substrates. Cell proliferation studies indicated that all samples had significantly lower cell doubling times than standard plastic monolayer (p < 0.001). Results confirm that graphene is not cytotoxic to DPSCs, which supports previous research.²⁷

SEM/EDX indicated that DPSCs grown on thick PB soft substrates appeared to have increased proliferation but limited biomineralization. In contrast, cells on the hardest substrates, 1:1 G-thin PB and thin PB, exhibited slower proliferation, but formed more calcium phosphate crystals, indicating greater biomineralization and osteogenic differentiation. The success of G-PB substrates in inducing osteogenic differentiation may be explained by the behavior of graphene itself. Graphene can influence cytoskeletal proteins, thus altering the differentiation of DPSCs through chemical and electrochemical means, such as hydrogen bonding with RGD peptides.^29,30 In addition, G-PB substrate stiffness may upregulate levels of alkaline phosphatase and osteocalcin, creating isometric tension in the DPSC actin network and resulting in greater crystal formation.³⁰ Overall, the proliferation results indicate that cells that undergo higher proliferation will undergo less crystal formation and osteogenic differentiation (and vice-versa).

The data presented here indicate that hard-soft intercalated substrates have the potential to enhance both proliferation and differentiation of DPSCs. G-PB substrates possess greater differentiation capabilities, whereas PS-PB substrates possess greater proliferative capabilities. Within graphene-based substrates, 1:1 G-thin PB induced the greatest biomineralization, performing better than various other substrates induced with DEX. This indicates that substrate stiffness is a potent stimulus that can serve as a promising alternative to biochemical factors like DEX.

Conclusion

The development of an ideal scaffold has been the focus of significant research in regenerative medicine. Altering the mechanical environment of the cell offers several advantages over current strategies, which are largely reliant on growth factors that can lead to acceleration of cancer metastasis. Within this study, the optimal scaffold for growth and differentiation of DPSCs was determined to be the 1:1 G-thin PB sample, which exhibited the greatest cell modulus, crystal deposition, and biomineralization. In addition, our study indicates two key relationships: one, the correlation between substrate and cell rigidity, and two, the tradeoff between scaffold-induced proliferation and scaffold-induced differentiation of cells, which depends on substrate characteristics. Further investigation of hard-soft intercalated substrates holds potential for developing safer and more cost-effective bone regeneration scaffolds.

References

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Developments in Gold Nanoparticles and Cancer Therapy

Catalyst Volume 9

Abstract

Nanotechnology has recently produced several breakthroughs in localized cancer therapy. Specifically, directing the accumulation of gold nanoparticles (GNPs) in cancerous tissue enables the targeted release of cytotoxic drugs and enhances the efficacy of established cancer therapy methods. This article will give a basic overview of the structure and design of GNPs, the role of GNPs in drug delivery and localized cancer therapy, and the challenges in developing and using GNPs for cancer treatment.

Introduction

Chemotherapy is currently the most broadly utilized method of treatment for most subtypes of cancer. However, cytotoxic chemotherapy drugs are limited by their lack of specificity; chemotherapeutic agents target all of the body’s most actively dividing cells, giving rise to a number of dangerous side effects.¹ GNPs have recently attracted interest due to their ability to act as localized cancer treatments—they offer a non-cytotoxic, versatile, specific targeting mechanism for cancer treatment and a high binding affinity for a wide variety of organic molecules.² Researchers have demonstrated the ability to chemically modify the surfaces of GNPs to induce binding to specific pharmaceutical agents, biomacromolecules, and malignant cell tissues. This allows GNPs to deliver therapeutic agents at tumor sites more precisely than standard intravenous chemotherapy can. GNPs also increase the efficiency of established cancer therapy methods, such as hyperthermia.³ This article will briefly cover the design and characteristics of GNPs, and then outline both the roles of GNPs in cancer therapy and the challenges in implementing GNP-based treatment options.

Design and Characteristics of Gold Nanoparticles

Nanoparticle Structure

To date, gold nanoparticles have been developed in several shapes and sizes.⁴ Although GNPs have also successfully been synthesized as rods, triangles, and hexagons, spherical GNPs have been demonstrated to be one of the most biocompatible nanoparticle models. GNP shape affects accumulation behavior in cells.⁴ A study by Tian et. al. found that hexagonal GNPs produced a greater rate of vesicular aggregation than both spherical and triangular GNPs.

Differences in GNP shape also cause variation in surface area and volume, which affects cellular uptake, biocompatibility, and therapeutic efficiency.⁴ For example, GNPs with greater surface area or more vertices possess enhanced cell binding capabilities but also heightened cell toxicity. Clearly, GNP nanoparticles must be designed with respect to their intended function.

Nanoparticle Surface Modification

In order to target specific cells or tissues, GNPs must undergo a ligand attachment process known as surface modification. The types of ligand particles attached to a GNP affect its overall behavior. For example, ligand particles consisting of inert molecular chains can stabilize nanoparticles against inefficient aggregation.⁵ Polyethylene glycol (PEG) is a hydrocarbon chain that stabilizes GNPs by repelling other molecules using steric effects; incoming molecules are unable to penetrate the PEG-modified surface of the GNPs.⁵ Certain ligand sequences can enable a GNP to strongly bind to a target molecule by molecular recognition, which is determined by geometric matching of the surfaces of the two molecules.⁵

Tumor cells often express more cell surface receptors than normal cells; targeting these receptors for drug delivery increases drug accumulation and therapeutic efficacy.⁶ However, the receptors on the surface of tumor cells must be exclusive to cancerous cells in order to optimize nanoparticle and drug targeting. For example, most tumor cells have integrin receptors.⁷ To target these residues, the surfaces of the therapeutic GNPs can be functionalized with the arginine-glycine-aspartic acid (RGD) sequence, which binds to key members within the integrin family.⁸ Successful targeting can lead to endocytosis and intracellular release of the therapeutic elements that the GNPs carry.

An important factor of GNP therapy is the efficient targeting and release of remedial agents at the designated cancerous site. There are two types of GNP targeting: passive and active. In passive targeting, nanoparticles accumulate at a specific site by physicochemical factors (e.g. size, molecular weight, and shape), extravasation, or pharmacological factors. Release can be triggered by internal factors such as pH changes or external stimuli such as application of light.² In active targeting, ligand molecules attached to the surface of a GNP render it capable of effectively delivering pharmaceutical agents and large biomacromolecules to specific cells in the body.

Gold Nanoparticles in Localized Cancer Therapy

Hyperthermia

Hyperthermia is a localized cancer therapy in which cancerous tissue is exposed to high temperatures to induce cell death. Placing gold nanoparticles at the site of therapy can improve the efficiency and effectiveness of hyperthermia, leading to lower levels of tumor growth. GNPs aggregated at cancerous tissues allow intense, localized increases in temperature that better induce cell death. In one study on mice, breast tumor tissue containing aggregated GNPs experienced a temperature increase 28°C higher than control breast tumor tissue when subjected to laser excitation.⁹ While the control tissue had recurring cancerous growth, the introduction of GNPs significantly increased the therapeutic temperature of the tumors and permanently damaged the cancerous tissue.

Organelle Targeting

GNPs are also capable of specifically targeting malfunctioning organelles in tumor cells, such as nuclei or mitochondria. The nucleus is an important target in localized cancer therapy since it controls the processes of cell growth, proliferation, and apoptosis, which are commonly defective in tumor cells. Accumulation of GNPs inside nuclei can disrupt faulty nuclear processes and eventually induce cell apoptosis. The structure of the GNP used to target the nucleus determines the final effect. For example, small spherical and “nanoflower”-shaped GNPs compromise nuclear functioning, but large GNPs do not.¹⁰

Dysfunctional mitochondria are also valuable targets in localized therapy as they control the energy supply of tumor cells and are key regulators of their apoptotic pathways.¹⁰ Specific organelle targeting causes internal cell damage to cancerous tissue only, sparing normal tissue from the damaging effects of therapeutic agents. This makes nuclear and mitochondrial targeting a desirable treatment option that merits further investigation.

Challenges of Gold Nanoparticles in Localized Cancer Therapy

Cellular Uptake

Significant difficulties have been encountered in engineering a viable method of cellular GNP uptake. Notably, GNPs must not only bind to a given cancer cell’s surface and undergo endocytosis into the cell, but they must also evade endosomes and lysosomes.¹⁰ These obstacles are present regardless of whether the GNPs are engineered to target specific organelles or release therapeutic agents inside cancerous cells. Recent research has demonstrated that GNPs can avoid digestion by being functionalized with certain surface groups, such as polyethylenimine, that allow them to escape endosomes and lysosomes.¹⁰

Toxic Effects on Local Tissue

The cytotoxic effects of GNPs on local cells and tissues remain poorly understood.¹¹ However, recent research developments have revealed a relationship between the shapes and sizes of GNPs and their cell toxicities. Larger GNPs have been found to be more cytotoxic than smaller ones.¹² Gold nanospheres were lethal at lower concentrations, while gold nanostars were less toxic.¹³ While different shapes and sizes of GNPs can be beneficial in various localized cancer therapies, GNPs must be optimized on an application-by-application basis with regard to their toxicity level.

Conclusion

Gold nanoparticles have emerged as viable agents for cancer therapy. GNPs are effective in targeting malignant cells specifically, making them less toxic to normal cells than traditional cancer therapies. By modifying their surfaces with different chemical groups, scientists can engineer GNPs to accumulate at specific tumor sites. The shape and size of a GNP also affect its behavior during targeting, accumulation, and cellular endocytosis. After accumulation, GNPs may be used to enhance the efficacy of established cancer therapies such as hyperthermia. Alternatively, GNPs can deliver chemotherapy drugs to tumor cells internally or target specific organelles inside the cell, such as the nucleus and the mitochondria.

Although some research has shown that GNPs themselves do not produce acute cytotoxicity in cells, other research has indicated that nanoparticle concentration, shape, and size may all affect cytotoxicity. Therefore, nanoparticle design should be optimized to increase cancerous cell death but limit cytotoxicity in nearby normal cells.

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Microglia: Gardeners with Guns

Catalyst Volume 9

If you ask anyone about the brain, their response will almost certainly involve neurons. Although neurons have been the stars of neuroscience for the past hundred years, the brain would be entirely dysfunctional if not for the variety of brain support cells, collectively known as glia.¹

Glial cells serve a variety of purposes in the central nervous system. Oligodendrocytes produce an insulating fatty-material called myelin, and astrocytes maintain electrical impulses in the neuronal network.¹ Perhaps the least glorious of glial functions are carried out by the microglia, which are the neurological equivalent of your household gardeners: pruning unwanted synapses and tending to the new ones. However, microglia are the first line of immune defense in the brain. From the brain’s humble beginnings as a mass of undifferentiated neurons to its affliction with the weeds of old age, microglia are tasked with neuronal maintenance and repair, meaning that deviation from their “just-right” activity can cause a variety of neural diseases. Too little activity, and one can be autistic or schizophrenic; too much activity, and one may be afflicted with Alzheimer’s or Parkinson’s. Given the large role these tiny cells play in brain protection, therapies that regulate microglial activation could be the key to curing a slew of neurological disorders.

Microglia respond to neural stress and injury through different mechanisms unique to their respective cell types: amoeboid phagocytic, resting ramified, and activated.² Amoeboid phagocytic glia act similarly to other scavengers and ingest large amounts of cellular debris in the developing brain during gestation.³ In postnatal development, these glia transform into resting ramified glia, which remain semi-dormant until their extended branches are activated by electrical signals from neurons or the presence of harmful substances.⁴ Activated microglia can secrete a variety of anti-inflammatory chemicals to prevent neurological problems, such as brains tumors and axonal injury.⁵ Microglia can also increase the permeability of the blood-brain barrier, allowing bodily immune cells to assist with brain immune defense.² A negative feedback mechanism in microglia regulates their own immune response as well as that of other helper immune cells.

In most pathologies, microglia experience a change in their normal activity caused by environmental factors.² Gliomas, or tumors in the neural glial tissues, are diseases that microglia should be able to handle. However, cells from the two microglial subcategories that migrate toward gliomal cells, M1 and M2, react differently in the gliomal microenvironment. M1 microglia promote tumor degradation by activating other immune cells and phagocytizing gliomal tumor cells. However, M2 microglia promote tumor growth by inhibiting proinflammatory cytokine activity and slowing immune cell responses.⁶ Cytokines are small proteins that aid cell communication and regulate cellular immune response.⁷ Additionally, tumor necrosis factor (TNF) stimulates inactivated microglial migration into the glioma, carving a pathway for glioma to migrate to other areas of the brain. Some gliomal therapies have focused on inhibiting the activity of M2 microglia. Various drug treatments that inhibit M2 activity have been shown to decrease gliomal proliferation in vivo. However, the success of these therapies should be treated with caution: gliomal immunosuppression both inactivates multiple immune responses outside of microglia and has the plasticity to circumvent anti-tumor therapies.⁶

Reduced microglial activity is related to a variety of neurodevelopmental disorders such as autism that demonstrate decreased connectivity in the brain.⁸ Microglia are responsible for forming mature spines and eliminating immature connections in the brain during post-natal development. This seems counterintuitive; how can decreasing in microglial activity, which causes less synaptic pruning, somehow cause less connectivity in the brain? Reduced microglial activity is actually preventing the brain from eliminating immature spine connections, which leads to fewer mature connections. Failing to eliminate immature connections physically hinders other synapses from forming multiple connections. Techniques that would increase microglial activity include increasing CR3/C3 pathway activity, which triggers synaptic pruning via an unknown mechanism.⁹ Although microglial therapies might not entirely eliminate autism, which acts through a variety of known and unknown neurological mechanisms, there is potential for ameliorating some symptoms.

Microglia often experience increased sensitivity in the aging brain caused by an increased expression of activation markers.¹⁰ This leads to several inflammatory neurological illnesses, including Alzheimer’s disease (AD). Microglia are once again found to play contradicting roles in the progression of Alzheimer’s; their activity is critical in producing neuroprotective anti-inflammatory cytokines, removing cell debris, and degrading amyloid-β protein, the main component of amyloid plaques that cause neurofibrillary tangles.¹⁰ Alternatively, activating microglia runs the risk of hyper-reactivity, which can cause extreme detriment to the central nervous system. Non-steroidal anti-inflammatory drugs (NSAIDs) have been shown to decrease the amount of activated microglia by 33% in non-AD patients. Treatment on microglial cultures increased amyloid-β phagocytosis and decreased inflammatory cytokine secretion. However, this treatment did not alter the microglial inflammatory activity in AD patients. The ideal microglial therapy for neuroinflammatory illnesses would result in the expression of only positive microglial activity, such as amyloid-β degradation, and the elimination of negative activity, such as pro-inflammatory secretion. One mechanism that increases pro-inflammatory secretion is amyloid-β binding to formyl peptide receptor (FPR) on microglia. Protein Annexin A1 (ANXA1) binding to FPR has been seen to inhibit interactions between amyloid-β and FPR, which decreases pro-inflammatory secretion.

Central nervous system pathology researchers often speculate as to how certain bacteria and viruses are able to enter the brain and consider mechanisms such as increase in blood-brain barrier permeability and chemical exchange through cerebrospinal fluid. However, the discovery of nervous system lymphatic vessels may put much of this speculation to rest and open up an entirely new venue of neuroimmunological research.¹¹ The interaction between microglial immune function and these lymphatic vessels could introduce treatments that recruit microglia to sites where bacterial and viral infections are introduced into the brain. Alternatively, therapies that increase bodily immune cell and microglial interactions by increasing the presence of bodily immune cells in the brain could boost the neural immune defense. Other approaches could involve introducing drugs that increase or decrease microglial-activity into more accessible lymphatic vessels elsewhere in the body for proactive treatment of neonatal brain diseases. Although we have made some steps towards curing brain diseases that involve microglial activity, coordinating these treatments with others that increase neural immune defenses has the potential to create effective treatment for those afflicted by devastating and currently incurable neurological diseases.

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