# Closed-loop firing rate regulation of two interacting excitatory and inhibitory neural populations of the basal ganglia

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## Abstract

This paper develops a new closed-loop firing rate regulation strategy for a population of neurons in the subthalamic nucleus, derived using a model-based analysis of the basal ganglia. The system is described using a firing rate model, in order to analyse the generation of beta-band oscillations. On this system, a proportional regulation of the firing rate reduces the gain of the subthalamo-pallidal loop in the parkinsonian case, thus impeding pathological oscillation generation. A filter with a well-chosen frequency is added to this proportional scheme, in order to avoid a potential instability of the feedback loop due to actuation and measurement delays. Our main result is a set of conditions on the parameters of the stimulation strategy that guarantee both its stability and a prescribed delay margin. A discussion on the applicability of the proposed method and a complete set of mathematical proofs is included.

### Keywords

Neural oscillations Firing rate models Time-delay systems Basal ganglia Parkinson’s disease Deep brain stimulation Closed-loop stimulation## 1 Introduction

Parkinson’s disease is a movement disorder associated with a loss of dopaminergic neurons in the substantia nigra, a nucleus located in the basal ganglia. In its initial stage, the disease’s symptoms can be relieved using dopaminergic medications. In more advanced stages, surgery is one of the alternatives proposed to some patients (Davie 2008). Even if ablative surgery of the subthalamic nucleus has been considered in the past (Bergman et al. 1990), it has now been replaced by deep brain stimulation (DBS), that is by the implantation of a lead that delivers electrical impulses (Benabid et al. 1991). This procedure has the advantage of being reversible (Limousin et al. 1995). Currently, there is no consensus on the mechanisms that explain the therapeutic effects of DBS (McIntyre et al. 2004a). Nevertheless, some clinicians seem to agree on the fact that the treatment is not always optimal and that there is still room for improvements (Kumar et al. 2003).

One of the directions proposed recently in order to optimize the efficiency of DBS is closed-loop stimulation (Rosin et al. 2011). Indeed, in its present form, DBS is an intrinsically open-loop procedure. The amplitude and frequency of stimulation are adjusted by the surgeon after the operation, but then these parameters remain fixed, with only sporadic adjustments during medical visits (Kumar et al. 2003). It is therefore logical to ask if the use of real-time measurements of brain activity might improve the efficiency of brain stimulation signals. Several works in the literature have explored this direction (for recent surveys, see, for example, Schiff 2010 and Carron et al. 2013). On the one hand, following the observation that in the parkinsonian state, some areas of the basal ganglia exhibit an increased synchronization in the beta-band (Nini et al. 1995; Hammond et al. 2007), some works focus on the desynchronizing aspects of stimulation (see, for example, Hauptmann et al. 2005a, b; Pyragas et al. 2007; Tukhlina et al. 2007; Omel’chenko et al. 2008; Lysyansky et al. 2011; Franci et al. 2011, and Montaseri et al. 2013). While, on the other hand, exploiting a detailed computational model of the basal ganglia such as (Rubin and Terman 2004), some other works try to reduce the impact of the pallidal output on the thalamo-cortical relay function (see, for example, Feng et al. 2007; Agarwal and Sarma 2010; Liu et al. 2010, 2011; Dunn and Lowery 2013). Nevertheless, on both sides, there seems to be a difficulty in reducing the gap between theoretical developments and experimentation.

Several problems might be at the origin of this difficulty. A first one is the microscopic nature of the considered models, which describe the behaviour of individual neurons rather than neural populations in their whole (Pascual et al. 2006). A second problem is the intermittent character of the basal ganglia dynamics (Dovzhenok et al. 2013). One should stress, moreover, that some of these works focus on an electrical stimulation and do not consider other possible ways of stimulating a neural population. In this paper, we consider a mesoscopic model in order to propose an alternative approach that takes into account the first problem, while remaining compatible with the modelling of some intermittent behaviours (see Sect. 5). In addition, in order to remain open to the possibilities offered by non-electric stimulations (see, for example, Strafella et al. 2004; Tufail et al. 2010; Han et al. 2011), we restrict ourselves to an abstract generic closed-loop stimulation.

A certain number of mesoscopic models have been proposed recently in the literature, in order to analyse the possible sources of pathological oscillations (see, for example, Leblois et al. 2006; Coombes and Laing 2009; Nevado-Holgado et al. 2010; Pavlides et al. 2012; Pasillas-Lépine 2013; Haidar et al. 2014). One of the motivations for the development of such models is the tight relationship between dopamine depletion, motor symptoms, and the enhanced oscillatory activity in the beta-band of parkinsonian patients (Jenkinson and Brown 2011). To our knowledge, this kind of mesoscopic models has never been used in order to design closed-loop stimulation strategies. In this paper, we aim at filling this gap. We rely on the basal ganglia model developed by Nevado-Holgado et al. (2010) and show that pathological oscillations can be reduced by decreasing artificially the gain of the subthalamo-pallidal feedback loop. We show that this can be achieved with a simple proportional feedback on the subthalamic nucleus. It can be shown, however, that this strategy is poorly robust with respect to actuation or measurement delays. We therefore complement this proportional closed-loop stimulation with a well-chosen low-pass filter and show that the resulting strategy can cope with these inherent delays.

## 2 Firing rate dynamics

*u*(

*t*) that represents the effect of a generic subthalamic stimulation.

Representing the stimulation of basal ganglia with such a model could be considered, of course, as an abusive oversimplification. Nevertheless, three points should be mentioned in order to understand why the analysis of this model might be interesting. First, it is well known that the basal ganglia have a somatotopic organization (Nambu 2011) and that pathological oscillations do not take place in the whole nuclei, but only in specific subareas (Zaidel et al. 2010). Therefore, even if the previous model cannot represent the whole basal ganglia, one might expect it to describe what happens in a given subregion of the dorsolateral STN, where both beta-oscillations occur and functional organization of neurons is homogeneous (due to somatotopy). Second, even if other basal ganglia models that take into account deep brain stimulation are available (see, for example, Rubin and Terman 2004; Modolo et al. 2008; Santaniello et al. 2011), their complexity makes it difficult to study analytically the stability of control laws that are applied to them. Considering a simplified model is thus a natural option in order to analyse what are the main effects of a closed-loop stimulation. Third, as already said, considering a generic stimulation in the above model has the advantage (at least in theory) to allow the potential consideration of non-electric stimulations. Even though this would also imply the consideration of more detailed models that take into account the specificities of each actuation mechanism in order to be fully analysed. In a few words, despite its simplicity, we expect that this generic stimulation could give some insights on what should be the effects of adding a controlled excitation or inhibition to the considered interconnected excitatory/inhibitory system.

Parameters of Nevado-Holgado’s model (2010)

Parameter | Value | Description |
---|---|---|

\(\delta _{sg}\) | \(6\,\hbox {ms}\) | Delay from STN to GPe |

\(\delta _{gs}\) | \(6\,\hbox {ms}\) | Delay from GPe to STN |

\(\delta _{gg}\) | \(4\,\hbox {ms}\) | Internal GPe delay |

\(\tau _{s}\) | \(6\,\hbox {ms}\) | STN time constant |

\(\tau _{g}\) | \(14\,\hbox {ms}\) | GPe time constant |

\(M_s\) | \(300\,\hbox {spk/s}\) | STN maximal firing rate |

\(B_s\) | \(17\,\hbox {spk/s}\) | Firing rate at rest for STN |

\(M_g\) | \(400\,\hbox {spk/s}\) | GPe maximal firing rate |

\(B_g\) | \(75\,\hbox {spk/s}\) | Firing rate at rest for GPe |

\(v_{s}\) | \(27\,\hbox {spk/s}\) | Cortical input to STN |

\(v_{g}\) | \(2\,\hbox {spk/s}\) | Striatal input to GPe |

Parameter | Healthy state | Diseased state |
---|---|---|

\(w_{sg}\) | 19.0 | 20.0 |

\(w_{gs}\) | 1.12 | 10.7 |

\(w_{gg}\) | 6.60 | 12.3 |

\(w_{cs}\) | 2.42 | 9.2 |

\(w_{xg}\) | 15.1 | 139.4 |

*i*to nucleus

*j*are defined by a linear parametrization

*K*is a parameter that describes the evolution of Parkinson’s disease and \(w^H_{ij}\) and \(w^D_{ij}\) are, respectively, the interconnection gains for the healthy and diseased states.

In the absence of stimulation, in the healthy case, the only equilibrium point of the model is locally asymptotically stable. As a consequence, no STN-GPe endogenous oscillations take place. This is not the case for the pathological case, where the stronger synaptic weights between STN and GPe result in an increase in the pallido-subthalamic loop gain that compromises stability. This generates a limit cycle (for details, see, for example, Nevado-Holgado et al. 2010; Pavlides et al. 2012; Pasillas-Lépine 2013), whose frequency stands in the beta-band, thus correlating with experimental observations (Hammond et al. 2007). It is therefore natural to explore the idea of using the exogenous stimulation signal *u*(*t*) in order to restore the system’s stability in the pathological situation.

## 3 Closed-loop stimulation

*u*(

*t*) a slightly delayed signal \(u(t-d)\) is applied to the system, the pathological oscillation persists and is even amplified, as it is shown in Fig. 2. Taking into account such issues is important, since delays are systematically present in both measurement and stimulation devices (see Sect. 5.3). The phenomenon observed in Fig. 2 should not be confused with the approach considered in Tukhlina et al. (2007). Indeed, while in both cases a proportional feedback and a delay are taken into account, the analysed models are not the same. In our case, it results from the Nyquist plot in Fig. 2 that there exists a delay \(d_0\) such that the feedback loop is stable for any \(d<d_0\) and unstable for any \(d>d_0\). This situation is completely different from the one analysed in Tukhlina et al. (2007), where the delay plays a positive role in the control law and contributes to its success.

## 4 Stability analysis

The aim of this section is to develop the mathematical arguments that lead to the stimulation scheme (5) proposed in the previous section. In particular, it is shown how the parameters of this scheme must be tuned in order to reduce the STN-GPe loop gain and avoid the generation of oscillations.

### 4.1 Error dynamics in the frequency domain

### 4.2 Reducing the gain of the main feedback loop

*H*entirely lies within the unit circle, which prevents the encirclement of the critical point. Since, moreover, we have

**Proposition 1**

Proposition 1 thus states that the gain of the transfer function \(H_s\) can be reduced arbitrarily close to \(\sigma _s^\star \) by sufficiently reducing the filter bandwidth of the closed-loop stimulation.

It is worth stressing that the expression of the function \(Q(\omega _p)\) is given explicitly in the proof, see Eqs. (21), (22) and (24), together with \(\omega _a\), the maximum value of \(\omega _p\) for which it holds, see Eq. (20).

The case of \(\Vert H_g \Vert _{\infty }\) is simpler than that of \(\Vert H_s \Vert _{\infty }\). Indeed, it is easy to find conditions for which this gain is bounded and does not depend on the stimulation parameters \(k_p\) and \(\omega _p\). This point is explained by the following result.

**Proposition 2**

As it can be seen in Eq. (7), the quantity \(\sigma _s^{\star }\) depends explicitly on \(k_p\) and \(\bar{x}_s\), but it does not depend on the filter frequency \(\omega _p\). Indeed, we have the following result, which exploits the sigmoid shape of the STN activation function and, in particular, its saturation for small and large firing rates.

**Proposition 3**

For every \(\epsilon >0 \), there exist a firing rate bound \(\bar{x}_\epsilon >0\) and a minimal gain \(k_\epsilon >0\) such that if \(k_p>k_\epsilon \), and \(\bar{x}_s\in \left( 0,\bar{x}_\epsilon \right) \cup \left( M_s-\bar{x}_\epsilon ,M_s\right) \), then \(\sigma _s^\star <\epsilon \).

In other words, by increasing the stimulation gain \(k_p\), and taking the reference \(\bar{x}_s\) close either to 0 or to \(M_s\) (the maximum value of the activation function \(F_s\)), we can reduce as we want the loop gain \(\Vert H\Vert _{\infty }\).

### 4.3 Stability of the internal feedback loops

The main consequence of the previous section is that the gain \(\Vert H\Vert _{\infty }\) can be made as small as we want. But in order to reach this goal, one has to increase the gain \(k_p\), which acts against the internal stability of the transfer function \(H_s\). Nevertheless, by reducing the filter’s frequency \(\omega _p\), one can always restore the stability of \(H_s\). Indeed, we have the following result.

**Proposition 4**

Consider the transfer function \(H_s\). For each delay \(d>0\), each firing rate reference \(\bar{x}_s\), and each gain \(k_p>0\), there exists a frequency \(\omega _b>0\) such that the transfer function \(H_{s}\) is stable for every filter frequency \(\omega _p>0\) such that \(\omega _p<\omega _b\).

For the internal stability of the transfer function \(H_g\), the analysis is simpler. Indeed, the condition \(w_{gg}\sigma _g<1\) of Proposition 2 implies the stability of \(H_g\), independently of the value of the internal delay \(\delta _{gg}\). This comes from the fact that under this condition, the Nyquist plot of \(H_g\) lies entirely within the unit circle, regardless of the value of \(\delta _{gg}\) (see, for example, Pasillas-Lépine 2013).

### 4.4 Robustness

The previous set of propositions define precisely how the stimulation parameters must be tuned. First, one has to compute a bound on \(\Vert H_g\Vert _{\infty }\) using Proposition 2. Second, one has to choose \(k_p\) and \(\bar{x}_s\), using Propositions 1 and 3, in order to have \(\Vert H_s\Vert _{\infty } < \left( 1 / (w_{sg} w_{gs}) \right) / \Vert H_g\Vert _{\infty }\). Third, the internal stability of \(H_s\) must be guaranteed by a sufficiently low value of the filter bandwidth \(\omega _p\), with the help of Proposition 4. This choice must also respect the constraint imposed by Proposition 1. In other words, the filter frequency must satisfy the constraint \(\omega _p<\min (\omega _a,\omega _b)\).

Nevertheless, when one has to choose concrete stimulation parameters, a critical issue is the question of robustness. Indeed, for a given set of model parameters, one can easily check what are the constraints imposed by the previous propositions. But, in a practical stimulation experiment, those parameters might change and be different from those of the nominal model. This point is not necessarily a problem in our approach, since the constraints imposed by the different propositions are inequalities. One can thus consider (among all possible values of the model parameters) the worst case and then choose the parameters of the control law in such a way that they respect the constraints imposed by the worst case. In other words, what we need in order to check our conditions is not the actual value of the different parameters, but some bounds on their possible values. In this context, the small-gain approach introduced in Sect. 4.2 plays a key role in what concerns robustness issues. Indeed, it automatically provides a design that is robust with respect to variations of the delays \(\delta _{sg}\) and \(\delta _{gs}\).

Of course, our method is probably over-conservative. Indeed, another approach would be to represent all the system’s uncertainties using an appropriate template in the frequency domain (Doyle et al. 1990, Chapt. 4.1). For the delay in the STN/GPe loop, it can easily be modelled as a multiplicative uncertainty. But it is not obvious for us how to find a weighting function (and an appropriate uncertainty model) that combines both the STN/GPe loop delay uncertainty and the actuator delay uncertainty in a tractable way. Moreover, one of our objectives is to propose explicit analytic constraints on the controller parameters that guarantee the closed-loop stability of the system. Even if it would be possible to formulate our robust stability problem in the form of an \(H_\infty \) problem, it is unclear (at least for us) how this problem could be solved analytically. Of course, the price to pay in order to obtain our explicit analytic constraints is some level of conservativeness. In contrast, we believe that an \(H_\infty \) approach might lead to less conservative constraints, but obtained after the numerical resolution of a set of LMI’s (see, for example, Doyle et al. 1989; Gahinet et al. 1994).

Another issue concerns our choice of considering in this paper only constant discrete delays. This choice was dictated by the desire to keep the mathematical complexity of the paper as reduced as possible and, additionally, by the fact that both axonal delays and electronic acquisition delays are usually assumed to be constant. For control theory techniques that are able to handle time-varying delays, the interested reader will find in Fridman and Shaked (2002) and He et al. (2007) two classical references that consider this problem (see also Seuret et al. (2013), for more recent techniques).

## 5 Discussion

In this paper, a strategy for disrupting sustained beta-band oscillations in the STN/GPe network has been described, using a closed-loop stimulation of the STN in the parkinsonian state. This approach is based on a mesoscopic model of the basal ganglia and takes into account both actuation and measurement delays. For that purpose, a filtered stimulation has been proposed with precise criteria for the choice of the filter bandwidth and of the reference STN firing rate. The encouraging theoretical results presented here need, however, to be confronted to experiments in order to evaluate their ability to handle some important physiological and technological questions. First, although evidence that increased motor symptoms are linked to increased beta frequencies in the STN (Hammond et al. 2007), the precise relation between beta-band oscillations and bradykinesia is still debated. Second, fine modelling of the specificity of the employed stimulation device, including limitations in terms of reactivity and maximum amplitude or frequency of the delivered stimulation signal, was neglected in the present analysis. Physiological limitations in the ability of the neuronal population to be influenced by the stimulation signal also need to be carefully addressed. Only animal experimentation will validate the efficacy of the proposed stimulation strategy. The aim of this section is thus to discuss the compatibility of our approach with an actual experimental setting in order to evaluate its potential significance.

### 5.1 Measuring firing rates

When recording electrophysiological activity of brain matter, most signals can be divided into two kinds of activity. On the one hand, after band-pass filtering between 500 and 2000 Hz, multiunit or single-unit activities can be recorded, typically with 800 k\(\Omega \) or 2 M\(\Omega \) electrodes, respectively, standing for spiking activity from a few axons in the vicinity of the electrode, typically 50–350 microns (Legatt et al. 1980). On the other hand, after low-pass filtering the acquired signal under 200 Hz, local field potentials can be extracted, standing for somato-dendritic activities of neuronal volumes larger than those in which multiunit activities can be recorded, especially when resorting to electrodes with impedances as low as 300 k\(\Omega \) (typically 0.5–3 mm in the vicinity of the electrode Mitzdorf 1987).

The method proposed in this paper resorts to a mesoscopic firing rate, defined for a functionally homogeneous population of neurons (Dayan and Abbott 2001). It is the average, over a specified time interval and over a certain number of neurons that are recorded, of all the spikes measured for this population of neurons. At least for *in vitro* preparations, there are examples in the literature where such measurements have been obtained in real time (see Wagenaar et al. 2005, for neuron cultures; and Carron et al. 2014, for rodent slices).

The mesoscopic firing rate could be experimentally determined in at least two ways in animal models of Parkinson’s disease. A first option would be to record inside the STN with multiplot electrodes, each plot having impedances around 800 k\(\Omega \). It could then be possible to record from several units in the vicinity of each plot and extrapolate the mesoscopic firing rate based on the recording of these neurons. This may be a good and directly testable option for laboratory experiments but would be less realistic for chronic recordings because of the mechanical instability of the recording setup. The second option for experimental translation of our theoretical work in the setting of chronic closed-loop stimulation would be to perform LFP recordings and then determine spiking activity of the area of neurons recorded. To this end, we may rely on the recent advances in the understanding of the link between single-unit activity and LFP (Teleńczuk and Destexhe 2014) and resort to identification techniques to extract spikes from LFP recordings such as the Hammerstein–Wiener model proposed by Michmizos et al. (2012). The advantage of resorting to LFP is that it enables screening the electrophysiological activity of a larger volume of neurons, and is less sensitive to mechanical instability and gliosis than multiunits recordings.

### 5.2 Which kind of stimulation?

*u*(

*t*), the sign of which accounts for this inhibitory or excitatory effect, and the amplitude of which depends on the amplitude and frequency of the illumination and on the number of recruited neurons (Han and Boyden 2007). Observe that in Fig. 4, the stimulation signal is negative at all times and thus may require, in order to be implemented, the use of inhibitory opsins (Tye and Deisseroth 2012).

Nevertheless, since electrical stimulation is the standard for clinical applications, it would be of interest to adapt the model to account for stimulations of this nature. We emphasize, however, that this is a challenging question in view of the complexity of the impact of electrical DBS on each neuron, which may vary according to the stimulated compartment (soma, dendrite, or axon) and to the orientation of the neuron (McIntyre et al. 2004b).

Other stimulation techniques, including Strafella et al. (2004), Tufail et al. (2010), or Han et al. (2011), could also be considered as possible candidates for experimental testing of the proposed stimulation strategy, provided a fine modelling of their dynamical impact on the targeted neuronal population.

### 5.3 Delays in the feedback loop

Two kinds of delays can be considered in the proposed STN closed-loop stimulation model: the acquisition and processing delay on the one hand, and the actuation delay on the other hand. Both these delays are embedded in the delay *d* considered in (6). The acquisition and processing delay combines the delay between the occurrence of a spike and its counting as an event by the electrophysiological recording system and the computation time required for the signal to be sent to the actuating system. It includes the time necessary to compute the firing rate (including filtering and possible spike sorting) and the resulting value of the control law. Nowadays, the order of magnitude of this process is of a few milliseconds. Regarding actuation delays, unlike electrical stimulation, optogenetics neuromodulation techniques offer in deep brain structures both millimetric-spatial and millisecond-time resolution for selective inhibition or excitation resorting to both inhibitory or excitatory photosensitive ionic channels (called opsins) and fibre-coupled lasers delivering the appropriate wavelength. For example, the photocurrent rise time for photostimulable inhibitory ionic channels such as ArchT is 7.4 ms (Zhang et al. 2007), a delay that is compatible with those considered in our approach. Recently, some attempts of disrupting neural oscillations using optogenetics (Han et al. 2009) have been carried out with success (Gradinaru et al. 2009).

### 5.4 Intermittence of pathological oscillations

Once the cortical and striatal inputs are fixed, the model we use here generates undamped oscillations in STN and GPe. This constitutes a simplified view of what is reported for in vitro (Plenz and Kital 1999) and in vivo (Park et al. 2010) experiments, in which these oscillations appear in a more transient way. This intermittence of pathological oscillations may compromise the efficacy of closed-loop stimulation policies, as reported in Dovzhenok et al. (2013).

A possible way to take into account this intermittence in our model is through the variation of cortical and striatal inputs. Such variations would modify the equilibrium of the STN-GPe network, which would in turn modify the value of the considered slopes of the activation functions (7). Thus, striatal and cortical inputs corresponding to lower slopes would not lead to oscillations, whereas those corresponding to higher slopes would favour oscillations onset. Such an analysis has been conducted in Pasillas-Lépine (2013).

We believe that the closed-loop stimulation policy proposed here is likely to cope with the oscillation intermittence produced by this kind of external inputs. The reason for this is that our stimulation policy essentially boils down to artificially decreasing the STN-GPe loop gain, as explained in Sect. 4.2. We expect this gain reduction to be beneficial in terms of oscillation generation during both intense busts of activity and more quiet periods. A simulation that illustrates this point is shown in Fig. 5.

### 5.5 Originality and potential significance

The first originality of our computational approach aiming at performing a stimulation disrupting oscillations in the beta-band frequency in the STN-GPe network is to consider a mesoscopic and not a microscopic model (see, for example, Hauptmann et al. 2005b; Franci et al. 2011; Rubin and Terman 2004; Agarwal and Sarma 2010; Liu et al. 2011). This mesoscopic approach may be more robust and realistic from an experimental point of view due to in vivo electrophysiological recordings technical requirements in rodents, non-human primates, or parkinsonian patients. The second originality is to build a method that is robust with respect to experimental constraints such as the delays occurring into a closed-loop system due to measurement, computational processing, and finally actuation.

From a more remote perspective, this work may provide some insights into how to design strategies for efficient and less energy-consuming clinical neuromodulation of parkinsonian patients. And, more generally, this stimulation strategy may be used to disrupt synchronization observed in physiological systems, may they be noticed in pathological conditions like epilepsy (Bragin et al. 2002) or schizophrenia (Gonzalez-Burgos and Lewis 2008), or in healthy conditions (Gray et al. 1989) enabling researchers to probe the role of synchronization in both healthy and pathological conditions.

## Notes

### Acknowledgments

This work was financially supported by the European Commission through the FP7 NoE HYCON2 and by the region Ile-de-France through the Neurosynch project (RTRA Digiteo). The work of the fourth author was supported by the Government of the Russian Federation (grant 074-U01).

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